Volume 12 Number 4 IJSPT by SPTS

IJSPT

INTERNATIONAL JOURNAL OF SPORTS PHYSICAL THERAPY

VOLUME TWELVE issue FOUR AUGUST 2017

IJSPT

INTERNATIONAL JOURNAL OF SPORTS PHYSICAL THERAPY

Editor in Chief Michael L. Voight, PT, DHSc, OCS, SCS, ATC, CSCS Belmont University Nashville, Tennessee – USA

Associate Editors: John Dewitt PT, DPT, SCS, ATC The Ohio State University Sports Medicine Columbus, Ohio - USA

Senior Associate Editor Barbara Hoogenboom, PT, EdD, SCS, ATC Grand Valley State University Grand Rapids, Michigan - USA

Terry Grindstaff PhD, PT, ATC Creighton University Omaha, Nebraska - USA

Manuscript Coordinator Ashley Campbell, PT, DPT, SCS, ATC Associate Editor, Thematic Issues: Robert Manske, PT, DPT, Med, SCS, ATC, CSCS Wichita State University Wichita, Kansas – USA

International Associate Editors: Mario Bizzini, PT, MSc Schulthess Clinic Zürich – Switzerland Henning Langberg, PT, PhD, MSc Institute of Sports Medicine Copenhagen – Denmark

Editorial Board: Scott Anderson, PT, Dip Sport PT Northgate Physical Therapy Regina, Saskatchewan – Canada

Mark S. De Carlo, PT, DPT, MHA, SCS, ATC Accelerated Rehabilitation Indianapolis, Indiana – USA

Lindsay Becker, PT, DPT, SCS, CSCS The Ohio State University Sportsmedicine Center Columbus, Ohio – USA

Todd S. Ellenbecker, DPT, SCS, OCS Physiotherapy Associates Scottsdale Sports Clinic Scottsdale, Arizona – USA

Barton Bishop, PT, DPT, SCS, CSCS Sport and Spine Rehab of Rockville Rockville, Maryland – USA

John A. Guido, Jr., PT, MHS, SCS, ATC, CSCS Ochsner Health Systems New Orleans, Louisiana – USA

Turner A. “TAB” Blackburn, Jr., MEd, PT, ATC Clemson Sports Medicine and Rehabilitation Manchester, Georgia – USA

Elizabeth L. Harrison, PT, PhD, Dip Sport PT University of Saskatchewan Saskatoon, Saskatchewan – Canada

Lori A Bolgla, PT, PhD, MAcc, ATC Augusta University Augusta, Georgia – USA

Walter L. Jenkins, PT, DHS, ATC East Carolina University Greenville, North Carolina - USA

Robert J. Butler, PT, DPT, PhD Duke University Durham, NC – USA

Daniel S. Lorenz, PT, DPT, ATC, CSCS Providence Medical Center Kansas City, Kansas - USA

Duane Button, PhD, CSEP-CEP Memorial University St. John’s, Newfoundland and Labrador – Canada

Terry Malone, PT, EdD, ATC, FAPTA University of Kentucky Lexington, Kentucky – USA

Rick Clark, PT, DScPT, CCCE Air Force Academy Colorado Springs, CO – USA

Peter J. McNair, PT, PhD Auckland University of Technology Auckland – New Zealand

George J. Davies, PT, DPT, SCS, ATC, FAPTA Armstrong Atlantic State University Savannah, Georgia – USA

EDITORIAL STAFF & BOARD

Phil Page, PT, PhD, ATC, CSCS The Hygenic Corporation Akron, Ohio – USA

Timothy Uhl, PT, PhD, ATC University of Kentucky Lexington, Kentucky – USA

Mark Paterno, PT, PhD, MBA, SCS, ATC Cincinnati Children’s Hospital Medical Center Cincinnati, Ohio – USA

Mark D. Weber, PT, PhD, SCS, ATC University of Mississippi Medical Center Jackson, Mississippi – USA

Charles E. Rainey, PT, DSc, DPT, MS, OCS, SCS, CSCS, FAAOMPT United States Public Health Service Springfield, Missouri - USA

Kevin Wilk, PT, DPT Champion Sports Medicine Birmingham, Alabama – USA

Michael P. Reiman, PT, DPT, OCS, SCS, ATC, FAAOMPT, CSCS Duke University School of Medicine Durham, North Carolina – USA Mark F. Reinking, PT, PhD, SCS, ATC Saint Louis University St. Louis, Missouri – USA Jill Robertson, PT, MSc (PT), Dip Manip PT Beaverbank Orthopaedic and Sport Physiotherapy Halifax, Nova Scotia – Canada Kevin Robinson, PT, DSc, OCS Belmont University Nashville, Tennessee – USA Barbara Sanders, PT, PhD, SCS, FAPTA Texas State University-San Marcos San Marcos, Texas – USA Teresa L. Schuemann, PT, DPT, SCS, ATC, CSCS Colorado Physical Therapy Specialists Fort Collins, Colorado – USA Brandon Schmitt, PT, DPT, ATC PRO Sports Physical Therapy of Westchester Scarsdale, New York - USA Patrick Sells, DA, ES Belmont University Nashville, Tennessee – USA Laurie Stickler, MSPT, OCS Grand Valley State University Grand Rapids, Michigan – USA Steven R. Tippett, PT, PhD, SCS, ATC Bradley University Peoria, Illinois – USA Timothy F. Tyler, PT, ATC NISMAT Lenox Hill Hospital New York, New York – USA

Erik Witvrouw, PT, PhD Ghent University Ghent – Belgium

international JOURNAL OF

IJSPT

SPORTS PHYSICAL THERAPY

I N T E R N AT I O N A L J O U R N A L OF SPORTS PHYSICAL THERAPY

SPORTS PHYSICAL THERAPY SECTION

Editorial Staff

Executive Committee

Michael L. Voight, PT, DHSc, OCS, SCS, ATC Editor-in-Chief

Walter L. Jenkins, PT, DHS, LATC, ATC President

Barbara Hoogenboom, PT, EdD, SCS, ATC Grand Valley State University Grand Rapids, Michigan - USA Senior Associate Editor

Blaise Williams, PT, PhD Vice President Mitchell Rauh, PT, PhD, MPH, FACSM Secretary

Robert Manske, PT, DPT, Med, SCS, ATC, CSCS Wichita State University Wichita, Kansas – USA Associate Editor, Thematic Issues

Bryan Heiderscheit, PT, PhD Treasurer Stacey J. Pagorek, PT, DPT, SCS, ATC Representative-At-Large

Associate Editors John Dewitt PT, DPT, SCS, ATC The Ohio State University Sports Medicine Columbus, Ohio - USA

Administration Mark S. De Carlo, PT, DPT, MHA, SCS, ATC Executive Director

Terry Grindstaff PhD, PT, ATC Creighton University Omaha, Nebraska - USA

Mary Wilkinson Director of Marketing/Webmaster Managing Editor, Publications

International Associate Editors Mario Bizzini, PT, MSc Schulthess Clinic Zürich – Switzerland

Contact Information P.O. Box 431 Zionsville, Indiana 46077 877.732.5009 Toll Free 317.669.8276 Fax www.spts.org

Henning Langberg, PT, PhD, MSc Institute of Sports Medicine Copenhagen – Denmark Ashley Campbell Manuscript Coordinator Mary Wilkinson Managing Editor

Advertising Sales The International Journal of Sports Physical Therapy accepts advertising. Email Mary Wilkinson, Marketing Director, at mwilkinson@spts.org or contact by phone at 317.501.0805.

IJSPT is a bimonthly publication, with release dates in February, April, June, August, October and December.

I N T E R N AT I O N A L J O U R N A L OF SPORTS PHYSICAL THERAPY is a publication of the Sports Physical Therapy Section of the American Physical Therapy Association. IJSPT is also an official journal of the International Federation of Sports Physical Therapy (IFSPT).

ISSN 2159-2896

IFSPT

TABLE OF CONTENTS VOLUME 12, NUMBER 4 Page Number

Article Title

ORIGINAL RESEARCH 512 The Dynamic Leap and Balance Test (DLBT): A Test-Retest Reliability Study. Authors: Jaffri AH, Newman TM, Smith BI, Miller SJ 520

Inter- and Intra-Rater Reliability of Performance Measures Collected with a Single-Camera Motion Analysis System. Authors: Bates NA, McPherson AL, Berry JD, Hewett TE

527

Validity of Athletic Task Performance Measures Collected with a Single-Camera Motion Analysis System as Compared to Standard Clinical Measurements Authors: McPherson AL, Berry JD, Bates, NA, Hewett TE

535

Peak Hip Muscle Torque Measurements Are Influenced by Sagittal Plane Hip Position Authors: Bazett-Jones DM, Tylinksi T, Krstic J, Stromquist A

543

Building A Better Gluteal Bridge: Electromyographic Analysis of Hip Muscle Activity During Modified Single-Leg Bridges. Authors: Lehecka BJ, Edwards M, Martin L, Porter K, Haverkamp R, Thatch K, Sack RJ, Hakansson NA

550

Effects Of A Band Loop On Lower Extremity Muscle Activity And Kinematics During The BarBell Squat Authors: Foley R, Bullbrook B, Button DC, Holmes MWR

560

A Biomechanical Comparison Among Three Kinds of Rebound-Type Jumps in Female Collegiate Athletes Authors: Nariai M, Yoshida N, Imai A, Ae K, Ogaki R, Suhara K, Shiraki H

569

Cadaveric Evaluation of the Lateral Anterior Drawer Test for Examining Posterior Cruciate Ligament Integrity Authors: Seeber GH, Wilhelm MP, Windisch G, Coriolano HJA, Matthijs OC, Sizer PS

581

Relationships Among Common Vision and Vestibular Tests in Healthy Recreational Athletes Authors: Heick JD, Bay C, Dompier TP, McLeod TCV

592

Comparison of a Head Mounted Impact Measurement Device to the Hybrid III Anthropomorphic Testing Device in a Controlled Laboratory Setting Authors: Schussler E, Stark D, Bolte JH, Kang YS, Onate JA

601

Transversus Abdominis Elasticity During Various Exercises: A Shear Wave Ultrasound Elastography Study Authors: Hirayama K, Akagi R, Moniwa Y, Okada J, Takahashi H

607

The Comparison of the Lumbar Multifidus Muscles Function between Gymnastic Athletes with SwayBack Posture and Normal Posture. Authors: Mahdavie E, Rezasoltani A, Simorgh L

616

Can Runners Perceive Changes In Heel Cushioning As The Shoe Ages With Increased Mileage? Authors: Cornwall MW, McPoil TG

625

There Are No Biomechanical Differences Between Groups Classified By The Functional Movement Screen.â&#x201E;˘ Authors: de Oliveira RR, Chaves SF, Lima YL, Bezerra A, Almeida GPL, Lima PO

634

Running Injury Development: The Attitudes of Middle- and Long-Distance Runners and Their Coaches. Authors: Johansen KK, Hulme A, Damsted C, Ramskov D, Nielsen RO

642

Cavitation Sounds During Cervicothoracic Spinal Manipulation. Authors: Dunning J, Mourad F, Zingoni A, Iorio R, Perreault T, Zacharko N, de las Penas CF, Butts R, Cleland J

CASE STUDIES 655 670

Return to Running Following a Knee Disarticulation Amputation: A Case Report Authors: Diebal-Lee AR, Kuenzi RS, Rabago CA The Effects of a Multimodal Rehabilitation Program on Pain, Kinesiophobia and Function in a Runner with Patellofemoral Pain: A Case Report Authors: Passigli A, Capacci P, Volpi E

CLINICAL COMMENTARY 683 A Four-Phase Physical Therapy Regimen for Returning Athletes to Sport Following Hip Arthroscopy for Femoroacetabular Impingement with Routine Capsular Closure. Authors: Kuhns BD, Weber AE, Batko B, Nho SJ, Stegemann C 697

Energy System Development and Load Management through the Rehabilitation and Return to Play Process. Authors: Morrison S, Ward P, duManoir GR

IJSPT

ORIGINAL RESEARCH

THE DYNAMIC LEAP AND BALANCE TEST (DLBT): A TEST-RETEST RELIABILITY STUDY Abbis H. Jaffri, PT, MS1,2 Thomas M. Newman, PhD, ATC2 Brent I. Smith, DHSc, ATC2 Sayers John Miller, PhD, PT, ATC2

ABSTRACT Background: There is a need for new clinical assessment tools to test dynamic balance during typical functional movements. Common methods for assessing dynamic balance, such as the Star Excursion Balance Test, which requires controlled movement of body segments over an unchanged base of support, may not be an adequate measure for testing typical functional movements that involve controlled movement of body segments along with a change in base of support. Purpose/hypothesis: The purpose of this study was to determine the reliability of the Dynamic Leap and Balance Test (DLBT) by assessing its test-retest reliability. It was hypothesized that there would be no statistically significant differences between testing days in time taken to complete the test. Study Design: Reliability study Methods: Thirty healthy college aged individuals participated in this study. Participants performed a series of leaps in a prescribed sequence, unique to the DLBT test. Time required by the participants to complete the 20-leap task was the dependent variable. Subjects leaped back and forth from peripheral to central targets alternating weight bearing from one leg to the other. Participants landed on the central target with the tested limb and were required to stabilize for two seconds before leaping to the next target. Stability was based upon qualitative measures similar to Balance Error Scoring System. Each assessment was comprised of three trials and performed on two days with a separation of at least six days. Results: Two-way mixed ANOVA was used to analyze the differences in time to complete the sequence between the three trial averages of the two testing sessions. Intraclass Correlation Coefficient (ICC3,1) was used to establish between session test-retest reliability of the test trial averages. Significance was set a priori at p â&#x2030;¤0.05. No significant differences (p >0.05) were detected between the two testing sessions. The ICC was 0.93 with a 95% confidence interval from 0.84 to 0.96. Conclusion: This test is a cost-effective, easy to administer and clinically relevant novel measure for assessing dynamic balance that has excellent test-retest reliability. Clinical relevance: As a new measure of dynamic balance, the DLBT has the potential to be a cost-effective, challenging and functional tool for clinicians. Keywords: clinical test, functional performance, postural control. Level of Evidence: 2b

Department of Kinesiology, University of Virginia, Charlottesville, VA 2 Department of Kinesiology, The Pennsylvania State University, University Park, PA The Pennsylvania State Universityâ&#x20AC;&#x2122;s Institutional Review Board approved this project on April 12, 2015; Study ID # STUDY00002306.

CORRESPONDING AUTHOR Abbis H. Jaffri Exercise and Sport Injury Laboratory University of Virginia 210 Emmet Street PO Box 400407 Charlottesville, VA 22904-4407 E-mail: ahj8uw@virginia.edu

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INTRODUCTION Balance is considered to be an important component of motor performance tasks. It is controlled by the central nervous system with the help of input from the visual, tactile, proprioceptive and vestibular systems.1,2 There are two main types of balance, static and dynamic. Static balance is defined as maintaining postural equilibrium while holding the body in a stationary position and dynamic balance is maintenance of postural equilibrium while parts of the body are moving.3 Gambetta and Gray4 refer to balance as the most important component in athletic ability because balance is involved in nearly every movement that is performed in daily life. There are a number of valid and reliable methods that are used in testing the static balance of a person such as quiet standing on a force plate in a laboratory setting or the Balance Error Scoring System (BESS) in a clinical setting.5-7 However, these measures of static balance may not provide relevant information about balance capabilities required to perform dynamic physical tasks.8 One test that is extensively used in clinical and research settings for dynamic balance assessment is the Star Excursion Balance Test (SEBT).5,6,9-12 Although the SEBT has been shown to be reliable and valid for identifying deficits in some musculoskeletal conditions, the movement tasks required during testing mimic only a limited number of activities involved in sports (e.g. ballet, gymnastics and ice-skating). Kinzey and Armstrong,13 reported that the SEBT may not be appropriate for the clinical assessment of dynamic balance. Moreover, it was suggested that the reaching movements performed in the SEBT are not normal movements performed by the lower limb in the activities of daily living.13 The authors recommended that normal functional movements (e.g. stair climbing, etc.) would be more appropriate for developing a dynamic stability test.13 Another test that has been used to measure dynamic balance is the Modified Bass Test which requires leaping between marks on the ground, and trying to maintain a balanced position for five seconds with each leap.1,3 Although this test requires base of support changes and alternate limb weight bearing, the hops require minimal effort and cannot be considered challenging for an active population.1,3 Moreover, the test has standard jump

distances that are not normalized to the leg length or height of the subjects. In laboratory settings, time-to-stabilization (TTS) assessments including single jump landing onto a force plate are being employed for testing dynamic balance.14 This testing method was able to demonstrate differences in dynamic balance between injured and uninjured populations.14 Time taken by the participant to stabilize themselves on one foot after landing on the force plate from a jump was used as the dependent measure. Healthy subjects took approximately two seconds to balance themselves while subjects with chronic ankle instability (CAI) took approximately three seconds to balance themselves.14 Unfortunately, force plates are neither prevalent, nor cost-effective in clinical settings. Based on the literature, accurate, cost-effective and efficient clinically relevant tests that measure dynamic balance abilities during functional tasks that require alternating limb weight bearing and base of support changes are not currently available. Hence, the Dynamic Leap and Balance Test (DLBT) which mimics the movement patterns commonly involved in activities (e.g. walking, running, cutting etc.) performed in daily life and sports requires the controlled movement of body segments over a base of support that is serially changing with alternating limb weight bearing. The DLBT is a low-cost clinical test based upon the concepts of previous balance tests such as the BESS, SEBT, TTS and the Modified Bass Test. The DLBT is a dynamic balance test that mimics normal activities of daily living and sport activities requiring serial changes in base of support, alternating limb weight bearing and a level of effort that should be challenging to an active population. The purpose of this study was to determine the reliability of the DLBT by assessing its test-retest reliability. It was hypothesized that there would be no statistically significant differences between testing days in time taken to complete the test. METHODS Participants Thirty (11 females, 19 males) healthy individuals, from the university community (age 24.0 Âą 3.1 years,

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height 172.4 ± 8.2 cm, mass 72.3 ± 14.2 kg) volunteered to participate in this study. Potential participants completed a self-report questionnaire regarding current and previous injury history. Participants were excluded if they reported current pain, numbness or paresthesia in the lower back or lower extremities, had a significant orthopedic injury of the lower back or lower extremity within the prior year (e.g. disc herniation, fracture, ligamentous sprain etc.), had a significant low back or lower extremity surgery within the prior year (e.g. ACL reconstruction, hip arthroscopy, lumbar laminectomy etc.), were currently under the care of a physician or seeking rehabilitation for a lower extremity or low back injury or pain, had a head trauma, concussion or were cognitively impaired within the prior six months, were currently experiencing any concussion like symptoms such as nausea, dizziness, headache etc., were currently experiencing balance problems, or were managing any neurological conditions e.g. stroke, Parkinson’s disease, etc. All participants meeting the criteria for inclusion in the study gave informed consent after reading and signing forms approved by The Pennsylvania State University Institutional Review Board and prior to participation. Included participants averaged 6.06 on the Tegner Activity Scale. Procedures Participants were tested on two separate days with a difference of at least six days between the first and second test days. Participants were tested by the same investigator on each of the two testing days. The testing sessions were similar on both the days. On Day 1, demographic data (age, sex and physical activity level) was collected and anthropometric measures (height, mass, limb length, foot length) were taken. Dominant leg length was measured from the anterior superior iliac spine to the apex of the medial malleolus and foot length was recorded as well as shoe size. Leg dominance was determined by asking the participant which leg they would prefer for kicking a ball. Participants then rated themselves on the Tegner Activity Level Scale.15 On both days, participants performed the DLBT. This test incorporated the directional layout of the SEBT.12,20 The pattern of the DLBT consists of 11

Figure 1. (a) DLBT pattern for right dominant limb. (b) DLBT pattern for left limb dominant. DLBT=Dynamic Leap and Balance Test; (1) Anterior short, (2) Anterior long, (3) Anteromedial short, (4) Anteromedial Long, (5) Medial short, (6) Medial Long, (7) Posteromedial short, (8) Posteromedial long (9) Posterior short (10) Posterior long, (C) Central Target.

total targets in the same positional directions as the medial half of the SEBT for each foot, including one central target and two targets along each of the five directions i.e. anterior, anteromedial, medial, posteromedial and posterior for both the right (Figure 1a) and left (Figure 1b) limb. The directional lines were placed on a floor using 1 ½ inch cloth athletic tape with 6-inch diameter cardboard circles used to mark the center and peripheral targets. The short and long target distances for each participant were

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normalized using the participant`s measured leg length and normative data for the SEBT.16 The proximal of the two peripheral targets was placed at 100% of the SEBT normative reach distance (expressed as a percentage of leg length) and the distal target was placed at 150% of the SEBT normative reach distance in each of the respective directions. The 150% target in each direction was included to make the task more difficult and encourage the participant to leap. Participants began at the center target of the testing matrix standing on their dominant limb with the nondominant limb foot next to the stance leg medial malleolus. The test was initiated by a verbal command of “Go” from the lead investigator. The participant then leaped from their dominant limb to a predetermined target landing on their non-dominant limb. The operational definition for a leap was “an acceleration or taking off from one limb and landing on the other limb”. Once on the peripheral target, the participant immediately leaped back to the central target, landing on their dominant limb and trying to attain and then maintain balance for two seconds. Attainment of balance was assessed using criteria similar to the modified Balance Error Scoring System (BESS) criteria (1) touching down with opposite foot, (2) excessive hip abduction, (3) out of testing position for more than two seconds and/or (4) step, stumble or fall).7 Once the investigator noted a restoration of balance for two seconds an audible command of “Go” was given to indicate the participant could leap to the next peripheral target. Participants continued this pattern of leaping and balancing for a total of 20 leaps (five directions and two distances in each direction). If the participant missed the target upon landing, they were instructed to reposition on the target as quickly as possible. All participants began with the anterior direction and moved in a clockwise (left leg dominant) or counterclockwise (right leg dominant) manner through all of the matrix directions, finishing after leaping from the posterior direction. In each direction, the participant leaped to the short target before the long target. Total time (seconds), to complete this task, was measured using a stop watch by the same investigator for each trial. The DLBT was verbally explained to the participants and one demonstration of the test was provided. Participants were instructed to complete the DLBT as quickly as possible. The participants were given three

practice trials before performing three timed trials. Two minutes of rest were provided between each of the practice and timed trials with five minutes of rest separating the practice and timed trials. Participants were allowed to wear their shoes and arm/hand position was unrestrained. The average time for the three timed trials was used for data analysis. Statistical Analysis A repeated measures two-way mixed ANOVA was used to calculate an Intraclass Correlation Coefficient (ICC)(3,1) to examine the test-retest reliability of the DLBT. Statistical analysis was performed using SPSS for Windows version 21.0 (SPSS Inc. Chicago, IL). Statistical significance was set a priori at p ≤ 0.05. RESULTS Time values were Day1 49.55±5.06 seconds, Day2 48.88±4.31 seconds and an average time of 49.21±4.61 seconds taken by participant on both testing days. There was no statistically significant difference between the three trial averages on Day 1 and Day 2 (p =0.054)). An ICC (3,1) of 0.93 with a 95% confidence interval from 0.84 to 0.96 indicates an excellent level of test-retest reliability was achieved with the DLBT (Table 1). The Standard Error of Measurement (SEM) is 4.67 seconds and the minimally detectable change (MDC) is 12.94 seconds. Figure 2 displays a scatter plot of the correlation between testing results of the DLBT on Day1 and Day2. DISCUSSION DLBT was created as a measure of the type of balance and dynamic joint stability capabilities required in most activities of daily living and sport activities. This novel test demonstrated excellent with in the tester test-retest reliability (ICC(3,1)=0.93) that compares favorably to other commonly used measures of static and dynamic balance. It has been reported, that the intratester reliability of the BESS battery of tests that consist of six different static stance positions ranged from an ICC of 0.50 to an ICC of 0.80 for the individual tasks and an ICC of 0.74 for total BESS score.17 Dynamic balance tests such as the Modified Bass Test1 and the Y Balance Test™, a variation of the SEBT, is associated with respective intra-rater reliability ICCs of 0.82 and 0.85 to 0.91.18

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Table 1. Intraclass correlation between times of Day 1 and Day 2 of Dynamic Leap and Balance Test* (DLBT). Intraclass Correlation ICC

.925

95 % Confidence Interval Upper bound .848

Lower bound .963

†Statistical Analysis p-value <0.01

* Values are the coefficient of reliability between the time taken on Day1 and on Day2 † Statistical Analysis performed using a one-way mixed ANOVA. Significance was set at p≤0.05

of support.11,19,20 These tasks involve challenging an individual’s limits of stability by moving their center of mass as close to their limits of stability as possible without losing balance stability.

Figure 2. Scatter plot showing correlation between the time taken by participants to complete the DLBT on Day1 with the time taken by participants on Day 2. Avg.D1T= Average time, in seconds, of three trials taken by participants on Day1 Avg.D2T= Average time, in seconds of three trials taken by participants on Day2 Pearson`s correlation coefﬁcient r = 0.936.

A review of available, clinically-relevant, static and dynamic balance tests revealed a paucity of tests designed to measure dynamic balance capabilities required in common functional activities. Quiet stance tasks are commonly used to measure static balance capabilities while the SEBT11,19 and Functional Reach Test20 are commonly used to measure dynamic balance capabilities. The goal of quiet stance tasks is to maintain a prescribed posture and move as little as possible. Quiet standing balance tasks typically challenge the ability of the individual to maintain their center of mass in a position well within their base of support. The goal of the SEBT involves moving lower body segments as far as possible away from a stable and unchanging base of support while the Functional Reach Test has the goal of moving the upper body segments as far as possible away from a stable and unchanging base

An additional goal of dynamic balance tasks is to move along a self-determined pathway with as little variability as possible. In reality, both static and dynamic tasks are dynamic in the sense that they invoke movement at lower extremity joints. Quiet stance tasks invoke movement around the ankle joint(s) while dynamic tasks require movement around multiple lower extremity joints. The goal of the assessment is the differentiating factor between static and dynamic tests. Previous research suggests that balance capabilities may be specific to the imposed challenge, with static balance tasks being less able to identify balance deficits related to athletic injuries than the dynamic balance tasks.21 This specificity may also be relevant when comparing dynamic balance tests. The reaching movements produced during the SEBT or Functional Reach Test may mimic activities performed in ice skating, gymnastics, dance or activities of daily living such as putting items into a cupboard, but do not represent the type of challenge presented to an individual’s postural equilibrium and dynamic joint stability during activities like walking, running, and cutting. Instead of controlling moving body segments over a stable base of support, these activities/ tasks require the individual to serially alternate their base of support from limb-to-limb requiring attainment of postural equilibrium with each change. It can be suggested that challenges to postural control may be unique to each type of task and therefore create the need for various dynamic balance tests that mimic specific activities. There is agreement with the assessment of Kinzey and Armstrong13 that the SEBT does not mimic common functional activities.

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Pionnier et al.22 did kinematic analysis on reaching tasks of SEBT. They found that there was systematic trend of lesser joint range of motion (ROM) of ankle, knee and hip joints in a CAI group.22 It is a possibility that limited ROM may be is a greater predictor for exhibiting lesser reach distances in CAI group rather than ability to maintain balance. However, when looking at the nature of the task of the DLBT, it appears that it is less likely to be limited by ROM which shows that the DLBT may be a more robust measure of dynamic balance. There are a number of hopping tasks that are used in assessment with a variety of lower extremity pathologies and procedures such as anterior cruciate ligament reconstruction (ACL-R). Hopping tasks are most commonly used to measure strength and power post ACL-R as compared to a task that is specifically geared towards measuring or quantifying dynamic balance. Myer et al.23,24 used various hopping tasks as a high-level power measure making them more of a strength measure than specifically a balance assessment tool. They can be helpful in giving some useful information regarding balance deficits but they are not specific balance assessment tools. Hopping tasks require propelling on and landing with the same limb which does not fully resemble functional activities like running, walking or cutting.24 However, the leaping task involving reciprocal lower extremity movement in the DLBT more closely mimics walking and running activities. The tuck jump is another task that is used in the ACL-R population for assessing patientâ&#x20AC;&#x2122;s landing mechanics in an effort to prevent further ACL injuries and assess the progress of rehabilitation.25 It can be a good measure to reveal patterns of movement in lower extremity that may result in recurrent ACL injury after reconstruction but it may not be an appropriate tool for specifically assessing dynamic balance. Similarly, the tasks like bilateral squatting or step downs are used to assess the muscular strength and joint range of motion in the lower extremity.26,27 However, neither of these tasks were originally designed to challenge the mechanisms that are involved in maintaining balance. The Vail Sports Test is another assessment test that involves functional activities against resistance.28 However, like other ACL-R return to play assessment tools this task focuses more on assessing the muscular

strength, power and endurance of the patient than balance.28 Patients are graded on the basis of the strength, power and endurance they demonstrate during this task.28 Although these strength, power and endurance assessment tools used with ACL-R patients are suggestive of the balance capabilities of a patient, they do not provide a specific assessment of the postural control system and are not designed to identify any postural control deficits. DLBT was created in an attempt to mimic the postural control challenges commonly encountered with functional activities such as walking, running and cutting.29 Balance requirements of common athletic activities such as running and cutting were assessed as a first step in the design of the DLBT. These tasks required the following: 1) an ability to maintain balance in single-legged stance, 2) an ability to change and reestablish a base of support from limb to limb without loss of stability, 3) incorporation of directional movements seen with running and cutting activities and, 4) incorporation of a physical effort challenge similar to walking, running and cutting. To assess an individualâ&#x20AC;&#x2122;s ability to balance on one leg and change base of support from one limb to the other, the individual was required to attain a stable, quiet stance posture for two seconds every time they leaped to the center circle of the grid. This component was based upon the results of previous landing studies that utilized time-to-stabilization as a measure of postural stability.14,29 It has been shown that healthy human participants take approximately two seconds to stabilize themselves on force plates during a single leg landing following a jumping task while participants with chronic ankle instability took longer.14 The leaping movements used in the current study resemble the single leg landing task of previous studies making the two second stability requirement clinically relevant. The BESS criteria were used for qualitatively judging errors in stability and a stop watch for calculating total time. Participants were given a verbal cue to continue after they stabilized for two seconds in the central target. One half of the grid pattern utilized in the SEBT was used to ensure directional movements were similar to those encountered in walking, running and cutting. Finally, short and long leap targets representing 100% and 150% of average leg length normalized SEBT reach distances for each direction

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were included to make the test more challenging. The longer target distances requires the individual to explosively propel their body mass toward the target while leaping by creating forces of acceleration and then slow down the movement of their body mass during landing through the creation of significant deceleration forces. Additionally, to and fro changes in position to the targets at alternating distances also creates a greater challenge for the motor control system. These challenges are more similar to those experienced in running and cutting and require greater physical exertion than those encountered in the SEBT. It is hoped that this greater physical challenge will provide a more sensitive differentiating factor in identifying balance deficits. A test of dynamic balance which more closely mimics the balance requirements of an individual’s intended activity participation may possess greater ability to identify persistent balance deficits after injury or prior to participation. In prior studies, the SEBT has been used to identify differences in balance abilities in injured and non-injured populations.19,30-32 It may be suggested that the DLBT will provide greater or additive differentiating power with regard to identifying dynamic balance deficits than the SEBT. The clinical utility of the DLBT, like the SEBT, would make it a valuable low cost tool for the clinical assessment of dynamic balance. Reviewing all these balance tests and functional assessment tools, it can be suggested that the DLBT comes out as an excellent addition to tools used for balance assessment. An effective progression hierarchy might be a BESS (static balance), SEBT (dynamic balance with unchanging BOS, limited by ROM) and the DLBT (dynamic balance with serial changes is BOS). There were associated limitations with this study. This study was performed with only healthy adults for the purpose of establishing test-retest reliability. The results of this study can therefore not be generalized to other populations. Further examination of the DLBT in patient populations suffering from musculoskeletal disorders is required to establish validity and reliability. CONCLUSIONS Dynamic balance assessment is an important measure of the postural stability that provides valuable

information about the balance deficits in injured populations. Unfortunately, there are very few clinically-relevant tests available to quantify dynamic balance. Furthermore, the tests available, such as the SEBT, do not appear to accurately mimic the dynamic balance requirements of common daily and sporting activities. The DLBT is a dynamic balance assessment tool that more closely mimics the challenges to dynamic stability encountered in common daily and sporting activities. The results of this study demonstrated excellent test-retest reliability for the DLBT. Further analysis of this testing procedure is required before its general effectiveness as a clinical measure of dynamic stability can be determined. REFERENCES: 1. Tsigilis N, Zachopoulou E, Mavridis T. Evaluation of the specificity of selected dynamic balance tests. Percept Mot Skills. 2001;92(3):827-833. 2. Wegener L, Kisner C, Nichols D. Static and dynamic balance responses in persons with bilateral knee osteoarthritis. J Orthop Sports Phys Ther. 1997;25(1):13-18. 3. Blackburn J, Prentice WE, Guskiewicz KM, Busby MA. Balance and joint stability: the relative contributions of proprioception and muscular strength. J Sport Rehabil. 2000;9(4):315-328. 4. Gambetta V, Gray G. Everything in Balance Train Cond. 1995;2(2):15-18. 5. Gribble PA, Hertel J, Denegar CR, Buckley WE. The effects of fatigue and chronic ankle instability on dynamic postural control. J Athl Train. 2004;39(4):321. 6. Herrington L, Hatcher J, Hatcher A, McNicholas M. A comparison of Star Excursion Balance Test reach distances between ACL deficient patients and asymptomatic controls. Knee. 2009;16(2):149-152. 7. Mulligan I, Boland M, Payette J. Prevalence of neurocognitive and balance deficits in collegiate football players without clinically diagnosed concussion. J Orthop Sports Phys Ther. 2012;42(7):625632. 8. Wikstrom EA, Tillman MD, Chmielewski TL, Borsa PA. Measurement and evaluation of dynamic joint stability of the knee and ankle after injury. Sports Med. 2006;36(5):393-410. 9. Filipa A, Byrnes R, Paterno MV, Myer GD, Hewett TE. Neuromuscular training improves performance on the Star Excursion Balance Test in young female athletes. J Orthop Sports Phys Ther. 2010;40(9):551558.

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10. Gribble P, Hertel J, Denegar C. Chronic ankle instability and fatigue create proximal joint alterations during performance of the Star Excursion Balance Test. Int J Sports Med. 2007;28(03):236-242. 11. Gribble PA, Hertel J, Plisky P. Using the Star Excursion Balance Test to assess dynamic posturalcontrol deﬁcits and outcomes in lower extremity injury: a literature and systematic review. J Athl Train. 2012;47(3):339-357. 12. Hertel J, Braham RA, Hale SA, Olmsted-Kramer LC. Simplifying the Star Excursion Balance Test: analyses of subjects with and without chronic ankle instability. J Orthop Sports Phys Ther. 2006;36(3):131-137. 13. Kinzey SJ, Armstrong CW. The reliability of the star-excursion test in assessing dynamic balance. J Orthop Sports Phys Ther. 1998;27(5):356-360. 14. Ross SE, Guskiewicz KM, Yu B. Single-leg jumplanding stabilization times in subjects with functionally unstable ankles. J Athl Train. 2005;40(4):298. 15. Hambly K. The use of the Tegner Activity Scale for articular cartilage repair of the knee: a systematic review. Knee Surg Sports Traumatol Arthrosc. 2011;19(4):604-614. 16. Gribble PA, Hertel J. Considerations for normalizing measures of the Star Excursion Balance Test. Meas Phys Educ Exerc Sci. 2003;7(2):89-100. 17. Finnoff JT, Peterson VJ, Hollman JH, Smith J. Intrarater and interrater reliability of the Balance Error Scoring System (BESS). PM R. 2009;1(1):50-54. 18. Plisky PJ, Gorman PP, Butler RJ, Kiesel KB, Underwood FB, Elkins B. The reliability of an instrumented device for measuring components of the star excursion balance test. N Am J Sports Phys Ther. 2009;4(2):92-99. 19. Plisky PJ, Rauh MJ, Kaminski TW, Underwood FB. Star Excursion Balance Test as a predictor of lower extremity injury in high school basketball players. J Orthop Sports Phys Ther. 2006;36(12):911-919. 20. Duncan PW, Weiner DK, Chandler J, Studenski S. Functional reach: a new clinical measure of balance. J Gerontol. 1990;45(6):M192-M197. 21. Cohen H, Blatchly CA, Gombash LL. A study of the clinical test of sensory interaction and balance. Phys Ther. 1993;73(6):346-351. 22. Pionnier R, Découfour N, Barbier F, Popineau C, Simoneau-Buessinger E. A new approach of the Star

Excursion Balance Test to assess dynamic postural control in people complaining from chronic ankle instability. Gait Posture. 2016;45:97-102. 23. Myer GD, Martin Jr L, Ford KR, et al. No association of time from surgery with functional deficits in athletes after anterior cruciate ligament reconstruction: evidence for objective return-to-sport criteria. Am J Sports Med. 2012;40(10):2256-2263. 24. Myer GD, Schmitt LC, Brent JL, et al. Utilization of modified NFL combine testing to identify functional deficits in athletes following ACL reconstruction. J Orthop Sports Phys Ther. 2011;41(6):377-387. 25. Myer GD, Ford KR, Hewett TE. Tuck jump assessment for reducing anterior cruciate ligament injury risk. Athl Ther Today. 2008;13(5):39-44. 26. Park K-M, Cynn H-S, Choung S-D. Musculoskeletal predictors of movement quality for the forward step-down test in asymptomatic women. J Orthop Sports Phys Ther. 2013;43(7):504-510. 27. Loudon JK, Wiesner D, Goist-Foley HL, Asjes C, Loudon KL. Intrarater reliability of functional performance tests for subjects with patellofemoral pain syndrome. J Athl Train. 2002;37(3):256. 28. Garrison JC, Shanley E, Thigpen C, Geary R, Osler M, DelGiorno J. The reliability of the vail sport test™ as a measure of physical performance following anterior cruciate ligament reconstruction. Int J Sports Phys Ther. 2012;7(1):20-30. 29. Ross SE, Guskiewicz KM. Examination of static and dynamic postural stability in individuals with functionally stable and unstable ankles. Clin J Sport Med. 2004;14:332-338. 30. Bressel E, Yonker JC, Kras J, Heath EM. Comparison of static and dynamic balance in female collegiate soccer, basketball, and gymnastics athletes. J Athl Train. 2007;42(1):42-46. 31. Doherty C, Bleakley C, Hertel J, Caulfield B, Ryan J, Delahunt E. Dynamic balance deficits 6 months following first-time acute lateral ankle sprain: a laboratory analysis. J Orthop Sports Phys Ther. 2015;45(8):626-633. 32. Olmsted LC, Carcia CR, Hertel J, Shultz SJ. Efficacy of the Star Excursion Balance Tests in detecting reach deficits in subjects with chronic ankle instability. J Athl Train. 2002;37(4):501-506.

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INTER- AND INTRA-RATER RELIABILITY OF PERFORMANCE MEASURES COLLECTED WITH A SINGLE-CAMERA MOTION ANALYSIS SYSTEM Nathanial A. Bates, PhD2,3 April L. McPherson1 John D. Berry2 Timothy E. Hewett, PhD2,3,4,5,6

ABSTRACT Background: Previous reliability investigations of single-camera three dimensional (3D) motion analysis systems have reported mixed results. Purpose: The purpose of the current study was to determine the intra- and inter-rater reliability of a single-camera 3D motion analysis system for subject standing height, vertical jump height, and broad jump length. Study Design: Experimental in vivo reliability study. Participants: Twelve subjects (age 20.6 ± 4.9 years) from a cohort that included high school to adult athletes who participated in sports at a recreational or competitive level entered and completed the study. Performance measurements were collected by a single-camera 3D motion analysis system and two human testers for standard clinical techniques. Inter- and intra-class correlation coefficients (ICC (2,k), ICC (2,1)) were determined. Result: Intra-tester and inter-tester reliability were excellent (ICC ≥ 0.935) for single-camera system measured variables. Singlecamera system measurements were slightly more reliable than clinical measurements for intra-tester ratings (ICC difference 0.020) for the standing broad jump. Single-camera system measurements were slightly less reliable than clinical measures for both intra- and inter-specimen standing height (mean ICC difference 0.003 and 0.043, respectively) and vertical jump height (mean ICC difference 0.017 and 0.036, respectively). Conclusions: The excellent reliability and previously demonstrated validity of the single-camera system along the anterior-posterior axis indicates that single-camera motion analysis may be a valid surrogate for clinically accepted manual measurements of performance in the horizontal plane. However, this single-camera 3D motion analysis system is likewise reliable, but inaccurate, for vertically oriented performance measurements. Level of Evidence: 2b Key words: Athletic performance; clinical motion analysis; Kinect™; reliability

Mayo Graduate School, Mayo Clinic, Rochester, MN, USA Department of Orthopedic Surgery, Mayo Clinic, Rochester, MN, USA 3 Department of Biomedical Engineering and Physiology, Mayo Clinic, Rochester, MN, USA 4 Sports Medicine Center, Mayo Clinic, Rochester, MN, USA 5 Department of Physical Medicine & Rehabilitation, Mayo Clinic, Rochester, MN, USA 6 Department of Biomedical Engineering, The Ohio State University, Columbus, OH, USA 2

Acknowledgements The authors acknowledge the support of the staff at the Materials Structural Testing CORE at Mayo Clinic, the Biomechanics Research Lab at Mayo Clinic, the Sports Health and Performance Institute at The Ohio State University. The authors acknowledge funding from NIH grants R01AR049735, R01AR056259, and R01AR055563 as well as the Mayo Clinic Foundation.

CORRESPONDING AUTHOR Nathaniel Bates Mayo Clinic 200 First St SW Rochester, MN 55902 Phone: (507)-538-6953 Fax: (507)-284-5392 E–mail: Bates.nathaniel@mayo.edu

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INTRODUCTION A variety of methodologies and technologies exist to analyze athletic performance as well as identify predictors of injury risk. Technological improvements have allowed in-depth analyses using three dimensional (3D) motion capture systems; however, these systems incur additional space, cost, and time requirements compared to manual measurements or two dimensional (2D) analyses.1 A critical requirement for the use of these technologies is the ability to reliably measure the subjects’ task performances both within and between testing sessions.2 Notably, between session reliability is necessary in order to monitor an athlete’s progression across a training or competitive season or through a rehabilitation program.2 If motion analysis technology is to be used to assess athlete and patient progression, then coaches and clinicians need to have confidence that the technology is outputting reliable data. Traditional 3D motion analysis systems exhibit good to excellent within- and between-session reliability.2,3 Previous investigations of reliability for single-camera 3D motion analysis systems utilizing the Microsoft Kinect™ sensor (Microsoft, Redmond, WA) have been reported. Inter-day reliability of the Kinect™ during gait assessment demonstrated modest to poor reliability for kinematic variables.4 Conversely, test-retest reliability recorded with the Kinect™ single-camera system for velocity and step length during gait has been shown to be excellent with intra-class correlation coefficients (ICC) ≥ 0.80.5 Similarly, a study that investigated peak joint angles during a squat with the Kinect™ system reported high between-trial reliability with ICCs > 0.90.6 These studies demonstrate both the uncertainty and potential of single-camera motion analysis systems to exhibit acceptable reliability for application in a clinical environment. The purpose of the current study was to determine the intra- and inter-rater reliability of the VirtuSense system (VirtuSense Technologies LLC., Peoria, IL), a single-camera 3D motion analysis system for subject standing height, vertical jump height, and broad jump length. It was hypothesized that both accepted clinical measurements and the VirtuSense system will demonstrate excellent intra- and inter-rater reliability for the selected performance tasks.

METHODS The current study was designed to investigate the intra- and inter-rater reliability of data outputs from a single-camera 3D motion analysis system for preprogrammed movements following a previously described methodology.7 Standing height, vertical jump assessment, and standing-broad jump distance were individually recorded with simultaneous measurement by VirtuSense Technologies and accepted clinical techniques performed by clinicians.8,9 The population cohort for this study included a sample of convenience comprised of high school to adult athletes who participated in sports at a recreational or competitive level. The same subject cohort was also used for a validation study. Exclusion criteria included history of lower extremity injury that prevented athletic participation within the last year. A total of twelve subjects (age 20.6 ± 4.9 years, nine male, three female) were recruited for participation. Six subjects were under the age of 18 and six subjects were over the age of 18. The Mayo Clinic Institutional Review Board approved the activities in this study and signed informed consent and assent (for participants under 18 years of age) documents were obtained prior to participation. The camera was placed on a level surface 91.4 cm above the ground, as recommended in the product user manual. Prior to testing, the subjects were allowed to perform their warm-up of choice. At a minimum, subjects were required to practice each performance task multiple times prior to data collection. Subjects performed each of the three tasks in front of the camera within the manual-specified capture volume. Task performance is explicitly described in the literature and briefly outlined here.7 For the standing height and vertical jump assessments, subjects stood with their heels on a mark 86.4 centimeters away from the camera. Vertical jump assessment was performed and peak height was recorded simultaneously by the VirtuSense and a Vertec measurement tool (Gill Athletics, Inc., Champaign, IL). Subjects were instructed to perform a maximal vertical jump and displace the highest possible vane on the Vertec device. Reach height for the Vertec was determined for each subject prior to performing the jumping task using a flat-footed, overlapped two-hand reach.8 For the standing broad jump assessment, subjects were

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positioned 3.96 meters away from the camera and instructed to jump towards the camera while maintaining balance upon landing.9 If the subject lost balance during the landing, they were instructed to repeat the task. Broad jump distance from a successful landing was recorded simultaneously by the single camera system and manually by a researcher who assessed the position of the big toe of the trailing foot against a tape measure affixed to the ground. The same researcher led every subject through each task three consecutive times. The duration of time elapsed between each trial of task performance was self-selected by each athlete. A second researcher then repeated the entire procedure with each subject. Subjects were afforded a brief self-selected rest period of no longer than 30 minutes between raters. All data for each individual subject were obtained during a single testing session. Performance of all athletic tasks was supervised by a certified athletic trainer with eleven years of experience. Discrete variables collected from each measurement task were evaluated for reliability with interclass correlation coefficients (ICC). Intra-rater (ICC (2,k)) and inter-rater (ICC (2,1)) correlations were examined separately for each of the three measured task (standing height, vertical jump height, and broad jump distance).10-12 Intra-rater reliability was determined using all three trials of a given task for each subject. Inter-rater reliability was determined using the mean of all three trial of a given task as collected separately by both of the individual raters. ICC calculations were conducted with MATLAB code (versions 2016a, Mathworks, Inc, Natick, MA) that has previously been verified against SPSS (version 20.0, IBM Corporation, Armonk, NY).3 ICC value ranges were classified based on previous literature where ICC < 0.4 was poor, 0.4 < ICC < 0.75 was fair-togood, and ICC > 0.75 was excellent.3,13,14 RESULTS Data collected by Tester 1 and Tester 2 for standing height, vertical jump height, and broad jump length are presented in Appendix 1. Both intra-tester and inter-tester reliability were excellent (ICC ≥ 0.958) for clinically measured variables (Table 1). Similarly, both intra-tester and inter-tester reliability were excellent (ICC ≥ 0.935) for single camera measured variables. The single-camera 3D system mea-

surements were slightly less reliable than clinical measures for both intra- and inter-specimen standing height (mean ICC difference 0.003 and 0.043, respectively) and vertical jump height (mean ICC difference 0.017 and 0.036, respectively). There was no difference between mean single-camera 3D system measurements and clinical ICC scores for broad jump length, while single-camera 3D system measurements were slightly more reliable than clinical measurements for intra-tester ratings (ICC difference 0.020). These minor differences did not affect reliability classification as every ICC calculated in this study demonstrated excellent reliability. DISCUSSION The objective of the current study was to determine the intra- and inter-rater reliability of a single-camera motion analysis system and accepted clinical measures for subject standing height, vertical jump height, and broad jump length performance. To the authors’ knowledge, this is the first study in the literature reporting the reliability of VirtuSense VirtuSports technologies. The results of this study support the hypothesis that both accepted clinical measurements and the VirtuSense system would demonstrate excellent intra- and inter-tester reliability. All of the ICC scores obtained for both clinically accepted and single-camera system measures were above 0.900 and classified as excellent. However, the single-camera 3D motion analysis system demonstrated lower intra- and inter-tester reliability scores than clinically measured values for standing height and vertical jump height. Conversely, the single-camera system exhibited better reliability than clinical measurements for the standing broad jump. Therefore, both the clinically accepted and singlecamera motion analysis measures should be reliably comparable across testers. The single-camera system exhibited excellent reliability in both the anterior-posterior and superior-inferior axes (Table 1). Although the anterior-posterior axis was slightly more reliable than the superior-inferior axis during dynamic tasks (ICC 0.977 versus 0.969 respectively), the differences were not a quantifiable classification difference. Previous investigation of the reliability of the Kinect™ system similarly reported greater reliability along the anterior-posterior axis compared to the superior-

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inferior axis.4 During a gait assessment, pelvis vertical range (a superior-inferior axis measurement) ICC was reported as 0.76 and 0.82 for comfortable and fast paces respectively, and step length ICC, an anterior-posterior axis measurement, was reported as 0.87 and 0.94 for the two pace conditions.4 This data corresponds to literature that demonstrated the Kinectâ&#x201E;˘-based VirtuSense system exhibited greater validity and accuracy in the anterior-posterior axis than the superior-inferior axis.7 Reliability along the superior-inferior axis was best for the standing height measurement, but this was to be expected since it was a static measurement. Previous literature reported greater error along the superior-inferior axis as well as an apparent saturation in output values.7 This may have caused the lower reliability of superior-inferior axis measurements reported in the current study, as increased potential for the introduction of measurement error can consequently decrease reliability. However, excellent reliability of measurements does not indicate excellent accuracy. A prior investigation that compared single-camera 3D system measurements to clinical measurements demonstrated significant differences in magnitude of values.7 This is reinforced by the slightly lower reliability scores along the super-inferior axis compared to the anterior-posterior axis. The single-camera system demonstrated excellent reliability for measuring standing height, vertical jump height, and broad jump length that is acceptable for clinical application; however magnitude of single-camera system outputs should be interpreted cautiously.7 A requirement for single-camera motion analysis systems to be an acceptable alternative measurement tool is that they should reflect comparable reliability to current measurement techniques. The clinical methods of measurement used in the current study have previously demonstrated excellent reliability.15,16 Clinical measurement values obtained in the current study exhibited excellent intra- and inter-tester reliability as expected. Mean intra- and inter- tester reliability in the current study for clinically measured vertical jump height was 0.986 and 0.971 respectively, which was greater than the intraand inter-session reliability previous reported for a sample of males (0.94 and 0.90 respectively).15 In

addition, mean intra-session reliability for clinically measured standing long jump was greater than previously reported (0.977 versus 0.930 respectively).16 The single-camera system was slightly more reliable than clinical measurements for the standing broad jump (inter-tester ICC 0.978 versus 0.958 respectively). Along the superior-inferior axis for the standing height and vertical jump height, clinical measures were slightly more reliable than the single-camera system (inter-tester ICC difference 0.043 and 0.036 respectively). These results confirm that the single-camera system demonstrates comparable reliability to current measurement techniques that is necessary for clinical application. However, the present data would also indicate that reliability and validity of single-camera motion analysis systems are not proportionally related, due to the inaccuracy in output magnitudes previously discussed.7 An additional consideration for implementing the single-camera system into the clinical environment is the ease of use. The two independent testers in the current study had different experience levels. Tester 1 had previous motion capture experience whereas tester 2 did not have previous motion capture experience. However, there were no differences in the classification of intra-ICC reliability scores for the VirtuSense measurements collected by each of the two testers. This finding suggests that the VirtuSense system can easily and reliably be adopted by a tester with limited previous experience. This study, as any study, is not without its limitations. All measurements were recorded during the same session and the greatest variation in reliability studies typically occurs between sessions.2,3,17 In a previous study where children performed lowerextremity reach tasks, excellent inter-rater reliability was reported within the same session, as well as excellent intra-rater reliability between sessions (ICCs > 0.90).17 However, the between session, single-rater reliability of this same cohort was 0.750.90, still rated as excellent but not as good as the within-session reliability. Furthermore, in the present study, the single-camera motion analysis system only recorded measurements along a single axis. Since joint angles are not reported, it remains unknown if the single-camera motion analysis system is an appropriate substitute for traditional 3D

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Table 1. Intra- and inter-rater reliability for clinically measured and VirtuSense performance measures.urements

Table 2. Intra- and inter-rater reliability for absolute differences in measurement between clinically measured and VirtuSense performance measurements.

motion analysis systems. Similarly, the reliability of this system in the analysis of total joint or whole body kinematics was not assessed in the current investigation. An additional limitation incurred by the single-camera motion analysis system is capture volume size. As previously reported in the literature, a single-camera system may not be able to accurately record an elite athleteâ&#x20AC;&#x2122;s full range of motion for certain athletic tasks, such as the vertical jump.7 This may limit the systems applicability to elite level athletes. The conclusions made in the present investigation are specific to tasks contained within the parameters of the user-manual-specified capture volume. CONCLUSION The single-camera VirtuSense motion analysis system demonstrated excellent intra- and inter-tester reliability comparable to clinically accepted manual measurements. However, excellent reliability is not a surrogate for validity and accuracy as demonstrated in prior literature.7 The excellent reliability and previously demonstrated validity of the single-camera system along the anterior-posterior axis indicates that

single-camera motion analysis may be a valid surrogate for clinically accepted manual measurements of performance in the horizontal plane. This claim is made with the caveat that the desired performance task fits comfortably within the capture volume of the single-camera motion analysis system. The system is likewise reliable, but not highly accurate, for vertically oriented performance measurements. REFERENCES 1. McLean SG, Walker K, Ford KR, et al. Evaluation of a two dimensional analysis method as a screening and evaluation tool for anterior cruciate ligament injury. Br J Sports Med. 2005;39(6):355-362. 2. Ford KR, Myer GD, Hewett TE. Reliability of landing 3D motion analysis: implications for longitudinal analyses. Med Sci Sports Exerc. 2007;39(11):2021-2028. 3. Myer GD, Wordeman SC, Sugimoto D, et al. Consistency of clinical biomechanical measures between three different institutions: Implications for multi-center biomechanical and epidemiological research. International Journal of Sports Physical Therapy. 2014;9(3):290-301. 4. Mentiplay BF, Perraton LG, Bower KJ, et al. Gait assessment using the Microsoft Xbox One Kinect: Concurrent validity and inter-day reliability of

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spatiotemporal and kinematic variables. J Biomech. 2015;48(10):2166-2170. Clark RA, Vernon S, Mentiplay BF, et al. Instrumenting gait assessment using the Kinect in people living with stroke: reliability and association with balance tests. Journal of neuroengineering and rehabilitation. 2015;12:15. Schmitz A, Ye M, Boggess G, et al. The measurement of in vivo joint angles during a squat using a single camera markerless motion capture system as compared to a marker based system. Gait Posture. 2015;41(2):694-698. McPherson AL, Berry JD, Bates NA, Hewett TE. Validity of athletic task performance measures collected with a single-camera motion analysis system as compared to standard clinical measurements. Int J Sports Phys Ther. 2017;12(4). Ferreira LC, Schilling BK, Weiss LW, et al. Reach height and jump displacement: implications for standardization of reach determination. J Strength Cond Res. 2010;24(6):1596-1601.

9. Myer GD, Schmitt LC, Brent JL, et al. Utilization of modiﬁed NFL combine testing to Identify functional deﬁcits in athletes following ACL reconstruction. J Orthop Sports Phys Ther. 2011;41(6):377-87. 10. Shrout P, Fleiss J. Intraclass correlations: Uses in assessing rater reliability. Psychol Bull. 1979;86(2):420 - 428.

11. Landis J, Koch G. The measurement of observer agreement for categorical data. Biometrics. 1977;33(1):159 - 174. 12. McGraw KO, Wong SP. Forming inferences about some intraclass correlation coefficients. Psychol Methods. 1996;1(1):30-46. 13. Fleiss JL. The design and analysis of clinical experiments. New York: Wiley; 1986. 14. Taylor-Haas JA, Hugentobler JA, DiCesare CA, et al. Reduced hip strength is associated with increased hip motion during running in young adult and adolescent male long-distance runners. IntJ Sports Phys Ther. 2014;9(4):456-467. 15. Nuzzo JL, Anning JH, Scharfenberg JM. The reliability of three devices used for measuring vertical jump height. J Strength Cond Res. 2011;25(9):2580-2590. 16. Moresi MP, Bradshaw EJ, Greene D, Naughton G. The assessment of adolescent female athletes using standing and reactive long jumps. Sports Biomech. 2011;10(2):73-84. 17. Faigenbaum AD, Myer GD, Fernandez IP, et al. Feasibility and reliability of dynamic postural control measures in children in first through fifth grades. Int J Sports Phys Ther. 2014;9(2):140-148.

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Appendix 1. Manually taken clinically accepted and VirtuSense measurements by tester 1 and tester 2

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ORIGINAL RESEARCH

VALIDITY OF ATHLETIC TASK PERFORMANCE MEASURES COLLECTED WITH A SINGLE-CAMERA MOTION ANALYSIS SYSTEM AS COMPARED TO STANDARD CLINICAL MEASUREMENTS April L. McPherson1 John D. Berry2 Nathanial A. Bates, PhD2,3 Timothy E. Hewett, PhD2,3,4,5,6

ABSTRACT Background: Previous investigations of single-camera 3D motion analysis camera systems validity have yielded mixed results for clinical applications. Purpose: The purpose of the current study was to determine the validity of a single-camera 3D motion analysis system for subject standing height, vertical jump height, and broad jump length. It was hypothesized that single-camera system values would demonstrate high correlation to the values obtained from accepted standard clinical measurements. Study Design: Experimental in vivo validation study. Methods: Twelve subjects (age 20.6 ± 4.9 years) from a cohort that included high school to adult athletes who participate in sports at a recreational or competitive level entered and completed the study. Performance measurements for standing height, vertical jump height, and broad jump length were measured with standard clinical measurements and a single-camera 3D motion system. Result: Single-camera system measurements were significantly different than clinical measures for standing height (p < 0.01) and vertical jump height (p < 0.01). There was no statistically significant difference between single-camera system measures and clinical measures for broad jump distance ( p > 0.07). The relative performance of subjects was highly correlated between single-camera and clinical measurements (r2> 0.80). Conclusions: Single-camera measurements lacked precision along the vertical axis of motion, but correlated well with clinically accepted measurements for standing height, broad jump length, and vertical jump height. The single-camera system may be capable of making accurate performance assessments in the horizontal plane, but should be limited to relative assessments along the vertical axis of motion. Additional refinement to increase the data reporting accuracy of the motion system along the vertical axis should be considered before relying on this single-camera 3D motion analysis system over clinical techniques to measure vertical jump and standing broad jump performances. Level of Evidence: 2b Key words: Athletic performance; Clinical motion analysis; Kinect™; Measurement validity; VirtuSense validation

CORRESPONDING AUTHOR Nathaniel Bates Mayo Clinic 200 First St SW Rochester, MN 55902 Phone: (507)-538-6953 Fax: (507)-284-5392 E–mail: Bates.nathaniel@mayo.edu

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INTRODUCTION Clinical evaluations of athletics tasks are often used as predictors of both athletic performance and injury risk potential. Clinicians, trainers, and coaches extensively rely on accepted techniques for manually measuring performance in tasks such as the vertical jump.1,2 Advances in technology have resulted in alternative methods that make a more in-depth analysis of task performance possible. Three dimensional (3D) analyses using multi-camera systems have been validated and provide additional information over two dimensional (2D) measurement systems.3 However, traditional 3D motion capture technology requires costly equipment and time-consuming data analyses, which include multiple cameras and processing software platforms.1,3,4 In order to minimize cost, space, and time requirements, single-camera 3D motion capture systems have been developed using Retro-Grate-Reflector (RGR) technology. Metria Innovation Tracking optical motion capture utilizes RGR technology.5 The current literature has demonstrated moderately good agreement in data between RGR systems and traditional 3D motion capture systems during athletic movements.5 The largest observed difference across three planes of motion was 1.2⬚ peak knee frontal plane angle, in which the RGR technology was 1.2⬚ less abducted than the multi-camera 3D motion analysis system during a single-leg land-and-cut maneuver. This 1.2⬚ difference represents approximately 14% of the average 8.7° change recorded by the 3D multicamera system in the frontal plane knee angle from touchdown to peak abduction.5 Additional literature demonstrates similar range of motion in the frontal plane during drop vertical jump tasks using a 3D multi-camera system.6 Single-camera Microsoft Kinect™-based systems (Microsoft, Redmond, WA) have also been investigated for use in the quantification of movement patterns in clinical settings. Comparison of hip and knee kinematics during jump landing and squatting movements between a traditional 3D motion analysis system and the Microsoft Kinect™ have demonstrated that the Kinect™ sensor may offer potential application as a real-time feedback tool.7 However, widespread adoption of the Kinect™ has not yet been

integrated into biomechanics due to limitations of the system. During a gait analysis, the accuracy of the Kinect™ sensor was deemed unacceptable for clinical application as its output kinematics exhibited low correlations and large errors in measurements when compared with a traditional 3D motion capture system.8 VirtuSense Technologies developed a single-camera 3D motion analysis system based on the moderate motion tracking success of previous single-camera sensors. This technology was designed to assess biomechanics and performance during dynamic tasks. However, prior to the introduction of VirtuSense into the clinical environment validation measurements are required. Previous investigation of the VirtuSense VirtuBalance tests demonstrated a lack of significant differences in outcome scores for Functional Reach Tests that were simultaneously collected by VirtuSense and a clinician, manually.9 Similarly, significant correlations between VirtuSense and clinically accepted measures for average postural sway led to acceptance of the VirtuSense as a valid balance assessment device.9 While this previous study contributes to validation of the VirtuSense system, the balance tests utilized elicit minimal perturbation from the participants.10,11 None of the balance tasks evaluated VirtuSense measurements along the vertical axis. For VirtuSense to be implemented as a clinical assessment tool for athletic performance, more rigorous movement tasks incorporating multiple axes of motion within the capture volume should be investigated. The objective of the current study was to determine the validity of a single-camera 3D motion analysis system for subject standing height, vertical jump height, and broad jump length. It was hypothesized that output from a single-camera motion analysis system would demonstrate high correlation to performance measurements obtained from standard clinical measurements used to clinically assess standing height, vertical jump height, and broad jump distance. METHODS The current study was designed to test the validity of data outputs from a VirtuSense motion analysis system (VirtuSense Technologies LLC., Peoria, IL)

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for movements pre-programmed in the VirtuSport system. Similar validity studies have previously been performed for the pre-programmed modified STAR and CTSIB balance tests within VirtuSense.9 Standing height, vertical jump assessment, and standing-broad jump distance were simultaneously recorded by VirtuSense Technologies and previously described clinical measurement techniques.2,12 The procedures in this study were approved by the Mayo Clinic Institutional Review Board. A subject cohort of convenience that included high school to young adult athletes who participate in sports at a recreational or competitive level were recruited for this investigation as well as a reliability study. Exclusion criteria included history of lower extremity injury that prevented athletic participation within the last year. Twelve subjects (age 20.6 Âą 4.9 years, nine male, three female) were recruited for participation. Signed informed consent and assent (for participants under 18 years of age) documents were obtained prior to participation.

ment. Participants were instructed to stand tall and face away from the device. Height was recorded to the nearest half centimeter by the researcher. Standing height was then measured virtually by the camera. The subject was positioned on a mark centered in front of the camera at a distance 86.4 cm away from the lens. Participants were instructed to stand tall facing the camera with their heels aligned to the designated mark on the ground. The VirtuSense output was recorded by the researcher. Proper collection technique of all standard clinical measurements was disseminated and overseen by a certified athletic trainer with over 11 years of experience. Vertical jump assessment height was performed and recorded simultaneously by the VirtuSense and a Vertec measurement tool (Gill Athletics, Inc., Champaign, IL). Reach height for the Vertec was determined using a flat-footed, overlapped two-hand reach.2 Subjects were instructed to perform a maximal vertical jump from a standing position without taking any lead-up approach. The tallest disrupted Vertec flag was recorded by the researcher. The output value recorded by the VirtuSense computer was also recorded by the researcher.

The camera was placed on the level surface of a 36 inch box during testing. This is within the parameters of recommended camera position as indicated by the camera user manual. Per the user manual, no device calibration is required if the camera is properly positioned on a level surface. For the standing height and vertical jump assessments, subjects stood with their heels on a mark 86.4 centimeters away from the camera. For the standing broad jump assessment, subjects started 396 cm away from the box and jumped towards the camera. The same researcher led every subject through each task three times. A second researcher then repeated the entire procedure a second time with each subject. Both researchers were trained in the use of VirtuSense technology by an individual who was formerly taught directly by VirtuSense personnel. In this manner the training afforded to the current researchers resembled the model of dissemination that typically occurs in the clinical environment. Tester 1 has approximately four years of experience in the use of motion capture technology, while Tester 2 was new to motion capture.

Standing broad jump assessment was also performed simultaneously by the VirtuSense and a tape measure. The tape measure was positioned on the floor directly below the camera and extended out away from the camera in a straight line. Participants were positioned along the tape measure facing the camera with their toes 3.96 meters distance away from the camera mount. This distance allowed the camera to maintain full view of the subject within the cameraâ&#x20AC;&#x2122;s capture volume. Participants were instructed to jump towards the camera while maintaining balance upon landing.12 If the subject lost balance during the landing, they were instructed to repeat the task. During successful trials, the researcher recorded the maximal toe distance from the initial starting position. Maximal toe distance was determined as the point of the toe of the trailing foot when landing was completed. Pilot testing determined that the furthest toe position was the most appropriate comparison to the VirtuSense measure.

Standing height was first recorded by a stadiometer with each subject standing erect with their back against the measuring pole. Participants did not remove their shoes for standing height measure-

Statistical analyses were performed using built-in functions in MATLAB (version 2015b, Mathworks, Inc., Natick, MA). For each subject an average of all three trials in each given task were calculated into a

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single data point. Data were checked for normal distribution. Statistical significance was set at p â&#x2030;¤ 0.05. Paired t-tests were used to determine if there were significant differences between the means of the single-camera 3D system values and the clinically accepted measurements within each task (standing height, vertical jump height, broad jump distance). Error was calculated as a percentage based on the difference between single-camera 3D system and clinical accepted measurements taken with respect to the clinically accepted measurements. Correlations were calculated using the built-in correlation function within MATLAB. Statistical significance was set at p < 0.05 and r-squared values were interpreted on a scale previously documented in the literature.13 RESULTS Subject performance demonstrated high r-squared correlations between VirtuSense and clinical measurements (r-squared > 0.80; Figure 1). As such, subjects who exhibited increased performance during clinical tests were also generally documented to have higher VirtuSense scores. Error magnitudes observed between the VirtuSense and clinical measurements for the vertical jump and broad jump tasks did not correlate significantly with subject height, vertical distance traversed, or horizontal distance traversed (p > 0.05, r-squared < 0.20; Figure 2).

There was no statistically significant difference in mean values for broad jump distance between the single-camera 3D measurements and accepted clinical measurements (p > 0.07). However, single-camera 3D measurements were significantly different than standard clinical measurements for standing height (p < 0.01) and vertical jump height (p < 0.01; Table 1). There were no statistically significant differences in mean values recorded between the two testers for either clinical or single-camera 3D measurements (p > 0.05). The average magnitude of error between the single-camera 3D and clinical measurements were 42.3% and 41.7% of the total vertical jump height for tester 1 and tester 2, respectively (Table 2). Magnitude of error for standing height was 3.7% and 3.1%, while error for broad jump was 1.3% and 0.7% of the total horizontal jump distance. For both testers, the magnitude of error exhibited between measurement techniques during vertical jump height was greater than each of the other tasks (p < 0.01). The magnitude of error for standing height was greater than the error for broad jump for both testers (p â&#x2030;¤ 0.01). DISCUSSION The objective of this study was to determine the validity of a single-camera 3D motion analysis system compared to standard clinical measurement techniques.

Figure 1. Correlations between VirtuSense and manually taken clinically accepted performance measurements. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 530

Figure 2. Correlations between manually taken clinically accepted performance measurements and VirtuSense error magnitudes.

Table 1. Mean values and standard deviations (cm) for clinically accepted and VirtuSense measurements

The only previous validation study of VirtuSense technologies has focused on the balance assessments.9 To the authorsâ&#x20AC;&#x2122; knowledge, this is the first study that examined the validity of VirtuSport tasks. The results of this study support the hypothesis that the singlecamera system outcomes would express strong r-squared correlation with clinically measured values. However, despite these correlations, the distances measured by the single-camera system and clinically accepted methods were significantly different during both the standing height and vertical jump height tasks. Such outcomes indicate that the VirtuSense can detect relative changes in distance measurements, but

the magnitude of those measurements may lack accuracy along the vertical axis. The single-camera system showed promise in reporting relative assessments of performance, as demonstrated by the strong correlations (Figure 1). Values for broad jump did not differ between the single-camera system and clinically accepted measurements. However, the average magnitudes of the single-camera system outcomes were significantly different than clinicallymeasured values recorded for vertical measures, which is of concern (Table 1). The magnitudes of the assessments by VirtuSense

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Table 2. Mean and standard deviation of the magnitudes of error (cm) between VirtuSense and clinical measurements

are not reliable along the superior-inferior axis, as both standing height and vertical jump assessment values were underreported by the single-camera system compared to clinically accepted measurements. Thus, a single-camera 3D system may not be an accurate system for precise vertical measurements, but it may be valid for anterior-posterior (broad jump) movements that are confined within the prescribed capture volume. In the present study, the available capture volume was optimized as the sensor was placed between 61.0 cm and 91.4 cm off the ground on a level surface and the subjects began at a distance greater than 152 cm away from the camera as per the developer instructions. In addition to the absence of statistical difference from clinically accepted measurements in broad jump values, the single-camera system measurements expressed a lower average magnitude of error along the anterior-posterior axis than the superiorinferior axis (Table 2). The magnitude of error compared to clinically accepted measurements along the anterior-posterior axis was 1.3% and 0.7% for tester 1 and tester 2, which equates to a disparity of 3.2 cm and 1.6 cm over the mean broad jump distance of 238 cm. Magnitudes of error for measures along the superior-inferior axis increased compared to the anterior-posterior axis errors. For standing height, the single-camera system underreported

subject height by an average of 3.7% (6.6 cm) and 3.1% (5.5 cm) for tester 1 and tester 2. Although the magnitude of error is low, standing height is a static measurement that should allow for precise measurements on a calibrated system. During dynamic tasks such as the vertical jump, the observed static error is likely to be amplified. The magnitude of error for the single-camera system values were 42.3% (24.2 cm) and 41.7% (24.0 cm) of total vertical jump height for each tester respectively. It is therefore unrealistic to anticipate accurate output from the singlecamera system along the superior-inferior axis of movement. The technology demonstrates potential application for accurately measuring movements along the anterior-posterior axis, but should be further investigated for movements along the superiorinferior axis. Along the superior-inferior axis of motion, the magnitude of error and restrictions of the capture volume are an issue relative to high-level athletes. Vertical jump performance at the National Football League Scouting Combine has a record of 114.3 cm and the National Basketball Association Combine record is 116.8 cm.14,15 In the present investigation, the greatest single-subject average vertical jump height was 70.7 cm. At that height, which is well below the professional combine averages, the singlecamera system appeared to experience an issue with

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saturation as all three jumps for the specified subject were recorded as exactly 46.0 cm. This represents an error difference of 24.7 cm despite a variation of only 3.8 cm in the clinical measurements. Similar situations where the ceiling of the capture volume may have been saturated were observed in 4 out of the 12 subjects tested in the present investigation. This circumstance will only be exacerbated with higherlevel athletes. Aside from reporting incorrect values, this issue of saturation is a problem because it can artificially limit the correlation between the singlecamera system values and clinical measurements. A lack of correlation due to capture volume saturation has the potential to further limit the capabilities of the single-camera system as it would inhibit clinical ability to assess relative comparisons of performance intra-athlete. The single-camera system did not exhibit apparent shortcomings relative to subject anthropometrics or task performance if the task was contained within the bounds of the capture volume. Standing height did not correlate to the magnitude of error in either performance task (Figure 2). Furthermore, overall performance in the vertical and broad jumps did not correlate to the magnitudes of error observed within that trial. Although the vertical jump assessment had the largest magnitude of error, it yielded the lowest standard deviation among the three tests. Therefore, the relative imprecise measurements along the superior-inferior axis may be attributed to a possible offset value. Data examination revealed that the single-camera system recorded value is an average of 24.1 cm less than the clinically recorded value. This may be due to the lack of a system calibration. The instruction manual states the sensor does not require calibration if placed on a level surface. Moreover, it may indicate a saturation of the single-camera system capture volume and inability of the system to accurately discern subject position relative to the starting position along the superiorinferior axis. As with any study, this study was not without limitations. It should be noted the apparent limitation in capture volume, which in the current study was determined to be about 396 cm along the anteriorposterior axis. In determination of the system setup, if the subject started further than 396 cm from

the camera, the system had difficulty with identification of the subject. Limitations of the single-camera system capture volume have also been discussed in the literature, as the technology was cited to function optimally at a range of 244 to 366 cm along the anterior-posterior axis.9 This restricted volume has the potential to prove a limitation when a single-camera system is used with high-level athletes. Standing broad jumps greater than 244 cm place athletes near the recommended bounds of the capture volume. Maximal broad jump recorded at the National Football League Scouting Combine 373 cm and would likely cause the single-camera system to lose track of the subject as they approach the limit of the measuring threshold.14 CONCLUSION Single-camera 3D motion analysis measurements lacked precision along the vertical axis of motion, but correlated well with clinically accepted measurements for standing height, broad jump length, and vertical jump height. The single-camera motion analysis system may be capable of making accurate performance assessments in the horizontal plane, but should be limited to relative assessments along the vertical axis of motion. The system has limited application to elite level athletes as their performance measures may stretch the recommended boundaries of the single-camera system capture volume. Due to the limitation presented, additional development work will be necessary for the VirtuSense to accurately and precisely replace accepted clinical measures for the assessment of athletic performance in the vertical jump and broad jump tasks. REFERENCES 1. Hewett TE, Ford KR, Xu YY, et al. Utilization of ACL injury biomechanical and neuromuscular risk proďŹ le analysis to determine the effectiveness of neuromuscular training. Am J Sports Med. 2016;44(12):3146-3151. 2. Ferreira LC, Schilling BK, Weiss LW, et al. Reach height and jump displacement: implications for standardization of reach determination. J Strength Cond Res. 2010;24(6):1596-1601. 3. Ford KR, Myer GD, Hewett TE. Reliability of landing 3D motion analysis: implications for longitudinal analyses. Med Sci Sports Exerc. 2007;39(11):2021-2028. 4. Hewett TE, Myer GD, Ford KR, et al. Biomechanical measures of neuromuscular control and valgus

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loading of the knee predict anterior cruciate ligament injury risk in female athletes: A Prospective Study. Am J Sports Med. 2005;33(4):492501. Weinhandl JT, Armstrong BS, Kusik TP, et al. Validation of a single camera three-dimensional motion tracking system. J Biomech. 2010;43(7):14371440. Bates NA, Ford KR, Myer GD, et al. Kinetic and kinematic differences between ﬁrst and second landings of a drop vertical jump task: Implications for injury risk assessments. Clin Biomech. 2013;28(4):459-66. Eltoukhy M, Kelly A, Kim CY, et al. Validation of the Microsoft Kinect(R) camera system for measurement of lower extremity jump landing and squatting kinematics. Sports Biomech. 2016;15(1):89-102. Pﬁster A, West AM, Bronner S, et al. Comparative abilities of Microsoft Kinect and Vicon 3D motion capture for gait analysis. J Med Eng Technol. 2014;38(5):274-280. Dodge K, Lynch R, Tippett SR. Reliability of real time motion analysis system VirtuBalance as

10.

11.

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compared to traditional measurement methods of functional reach and postural sway. J Stud Phys Ther Res. 2015;8(5). Duncan PW, Weiner DK, Chandler J, et al. Functional reach: a new clinical measure of balance. J Gerontol. 1990;45(6):M192-197. Loughran S, Tennant N, Kishore A, et al. Interobserver reliability in evaluating postural stability between clinicians and posturography. Clin Otolaryngol. 2005;30(3):255-257. Myer GD, Schmitt LC, Brent JL, et al. Utilization of modified NFL combine testing to identify functional deficits in athletes following ACL reconstruction. J Orthop Sports Phys Ther. 2011;41(6):377-87. Shrout PE, Fleiss JL. Intraclass correlations: uses in assessing rater reliability. Psychologicial Bulletin. 1979;86(2):420-428. NFL Combine Top Performers. http://www.nfl.com/ combine/top-performers. NBA Pre-Draft Measurements. http://www. draftexpress.com

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IJSPT

ORIGINAL RESEARCH

PEAK HIP MUSCLE TORQUE MEASUREMENTS ARE INFLUENCED BY SAGITTAL PLANE HIP POSITION David M. Bazett-Jones, PhD, LAT, ATC, CSCS1 Tyler Tylinksi, PT, DPT1 Jelena Krstic, PT, DPT1 Abigail Stromquist, PT, DPT1 Jay Sparks, PT, DPT1

ABSTRACT Background: An optimal position for strength testing of the hip musculature has not been identified. However, sagittal plane hip position during testing has been shown to influence hip external rotation strength. Hypothesis/Purpose: The purpose of this study was to compare hip extension, external rotation, and abduction isometric torque at positions with differing degrees of hip flexion using a handheld dynamometer. Study Design: A cross-sectional laboratory study. Methods: Twenty-nine healthy and physically active females participated in this study. Peak isometric contractions were measured with a handheld dynamometer secured with a non-elastic strap and then converted to torque using segment lengths. Hip external rotation and extension were tested at 0 , 30 , and 90 of hip flexion. Hip abduction was tested at 0 and 30 of hip flexion and 5 of extension. Testing was randomized and counterbalanced. Repeated measures ANOVAs with Sidak’s test for multiple comparisons were used for statistical analysis. Significance was set at p<0.05. Results: Significant main effects were found for hip extension (p<0.001) and external rotation (p<0.027), but not for abduction (p=0.085). Pairwise comparisons showed significant differences between all three testing positions for hip extension torque (0 v30 : p<0.001, 0 v90 : p<0.001, 30 v90 : p=0.002). Extension torque was highest in 90 of flexion (1.43±0.50 Nm/kg*m) and lowest in 0 of flexion (0.83±0.30 Nm/kg*m). Comparisons of hip external rotation torque tested at 0 v90 (p=0.096) and 30 v90 (p=0.080) were not significantly different but did have medium effect sizes. External rotation torque was highest in 90 of flexion (0.29±0.13 Nm/kg*m). Conclusions: Direct comparisons of torque values of hip extension and external rotation tested at different sagittal plane positions should be cautioned due to differences. Hip extension and external rotation should be measured in consistent sagittal plane positions across examiners and testing sessions. Test position will be dependent upon the goals of strength testing. Level of Evidence: 2b Key words: Dynamometer, isometric force, hip joint, muscle strength

Department of Physical Therapy, Carroll University, Waukesha, WI, USA

CORRESPONDING AUTHOR David M. Bazett-Jones Carroll University 100 N. East Ave Waukesha, WI USA E–mail: dbazettj@carrollu.edu

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INTRODUCTION Hip strength has been identified as an important variable in multiple lower extremity pathologies. Weakness of the hip abductors (ABD) has been found following acute ankle sprains1 and in those with patellofemoral pain,2 patellofemoral osteoarthritis,3 exertional medial tibial pain,4 and chronic hip pain.5 Hip external rotation (ER) strength has also been found to be reduced in individuals with patellofemoral pain2 and iliotibial band syndrome,6 and those sustaining anterior cruciate ligament tears.7 Hip extension (EXT) strength has been found to be reduced following knee surgery,8 in those with chronic low back pain,9 and in those with lower extremity injury,9 specifically patellofemoral pain.10 Standardized testing of hip strength is vital for assessing, treating, and preventing multiple trunk and lower extremity pathologies. The optimal hip strength testing position has not been identified in the in the published research, which results in a great amount of variability between studies with regard to sagittal plane hip positioning. Studies assessing hip ABD strength have performed testing in positions of 0 of hip flexion (i.e. neutral)3,10,11 and varying degrees of hip extension.12,13 Hip EXT strength has been measured at 0 (i.e. neutral),8,9,11 30 ,10 and at 90 of hip flexion.14,15 Testing of hip ER has been measured at 0 of hip flexion,16,17 but is most commonly measured at 90 of hip flexion (i.e. seated).3,5,11 Given the variability in the literature, research is necessary to determine whether hip muscle torque production is different among the various sagittal plane hip positions used during testing. Differences would limit the ability to perform comparisons across studies using differing methodologies. Indeed, it is reasonable to suspect differences among testing positions would exist as the moment arms of muscles around the hip have been shown to change based on hip flexion position.18 A change in moment arms can result in a change in contributions by individual muscles and can influence torque production when tested in varying sagittal plane hip positions. For example, as the hip flexes, the moment arm of the ER muscles decreases,18 meaning that a flexed hip position could result in lower ER torque values. However, previous studies19,20 have reported that

hip ER torque values are greater at 90 than 0 of hip flexion. These authors19,20 have attributed these differences in ER torque to length-tension changes with changes in hip flexion position. Others21,22 have reported no difference in hip ER strength between positions of varying hip flexion. While hip ER strength has been previously investigated, the influence of sagittal plane hip position on EXT strength has had limited investigation. Jensen et al23 described peak EXT values occurring at 40 of hip flexion with values approaching zero at 80 of hip flexion and 15 of hip extension (past zero). Clark24 reported peak hip EXT values at full hip flexion (140 ) and decreasing values as the hip was extended to 10 of hip flexion. As the hip flexes, the moment arm of the gluteus maximus is reduced and the moment arms of the hip adductors and hamstrings are increased.25 It is likely that the lengthtension relationship of these muscles change as well as the hip is flexed. Given that hip strength has been tested in many different positions in the literature, it is important to understand differences in hip EXT strength due to testing position in order to ensure its appropriate and consistent measurement. Hip ABD strength has been tested in both 0 of hip flexion and in hip extension (5-10 ). As the hip flexes, the tensor fasciae latae might have a greater contribution to hip ABD. As the hip flexes, gluteus medius and minimus moment arms decrease and the tensor fasciae latae moment arm is maintained.25 Increased hip flexion may also explain increased tensor fasciae latae activity during clam exercises.26,27 However, the authors were unable to find any research comparing sagittal plane hip position on hip ABD strength testing. Previous authors19,20,23,24 have utilized isokinetic dynamometers and complex experimental devices to investigate the influence of sagittal plane hip position on hip strength. These devices are large, expensive, and may not be readily available or practical for clinical use or pre-participation screenings. Hand-held dynamometers (HHD) have been used extensively in the literature.3,5,6,11,13,16,17,28-30 However, further research is needed on whether differences in hip strength are present when using a HHD in different positions of hip flexion. Therefore, the purpose

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of this study was to compare hip EXT, ER, and ABD isometric torque at positions with differing degrees of hip flexion using a HHD. It was hypothesized that an increased hip flexion position would influence all strength tests. Specifically, that 1) hip EXT strength would be greatest at 90 and least at 0 of hip flexion, 2) hip ER would be greatest at 90 and least at 0 of hip flexion, and 3) hip ABD would be greatest at 30 of hip flexion and least at 5 of hip extension. METHODS Participants were recruited from a small university and it surrounding communities in the Midwest United States. Females (18-45 years of age) who were physically active and healthy were included in this study. Participants were considered physically active if they performed more than 30 minutes of physical activity three or more days/week. A self-report medical history and physical activity questionnaire was used to determine eligibility. Participants were excluded if they reported any back or lower extremity musculoskeletal pain, injury, or conditions in the prior 12 months, pregnancy, instability of any lower extremity joint, and head injury or neurological disorder that would influence testing. This study was approved by the Carroll University Institutional Review Board (#14-003) and participants were informed of study procedures and signed a written consent form prior to initiating study procedures. Thirty healthy, physically active women volunteered to participate in this study (mean±SD: height=1.63±0.07 m, weight=61.1±10.6 kg, age=21.1±3.3 years). Sample size for the strength variables was determined utilizing female ER torque data from Hoglund et al.19 An alpha value of 0.05, beta value of 0.20, and an effect size of 0.33 were used and it was determined that 17 participants were necessary for this study. Since ABD and EXT had not been previously tested, a sample size of 30 was utilized in order ensure adequate power for this study and to account for any attrition. One participant did not return for testing, resulting in 29 total participants for this study. The testing session consisted of height and weight measurement, followed by limb segment length and strength assessments. The tested limb was randomly selected (i.e. right or left) by coin toss. Limb segment lengths were measured to represent moment arms

for torque measurement calculations. Femur segment length was determined by measuring the distance between the greater trochanter and the lateral epicondyle. Leg segment length was determined by measuring the distance between the medial epicondyle and medial malleolus. These measurements were taken with a cloth tape measure as previously reported by Bolgla et al.30 Strength of the hip ABD, ER, and EXT was measured using maximal voluntary isometric contractions against a HHD (Model 01136, Lafayette Instruments, Lafayette, IN, USA) secured by non-elastic straps. Strength tests using similar HHD methods have demonstrated excellent reliability.28,29,31 For this test, participants were instructed to contract maximally for three to five seconds and were given standardized verbal encouragement. Participants were given one practice trial and three maximal trials with one minute of rest between trials to provide adequate rest. Hip ABD was tested in side-lying with the hip in 0 and 30 of hip flexion, and 5 of hip extension (Figure 1, middle row), 0 of hip abduction and rotation, and full knee extension (0 ). The non-tested limb was placed in hip and knee flexion to provide stability with a pillow between the thighs to maintain a neutral position. The HHD was placed just proximal to the lateral epicondyle of the knee, while an extra stabilizing strap was placed over the iliac crest to prevent extraneous pelvic movements. Hip ER was tested in prone with the hip in 0 and 30 of flexion (Figure 1, top row), and seated with the hip in 90 of flexion (Figure 1, top row). The tested limb was positioned in 90 of knee flexion and 0 of hip abduction and rotation. The HHD was placed just proximal to the medial malleolus, while a stabilizing strap was placed over the thighs to prevent extraneous movements. Hip EXT was tested in prone with the hip in 0 , 30 , and 90 of flexion (Figure 1, bottom row) while keeping the knee in 90 of flexion and the hip in 0 of rotation. The HHD was placed just proximal to the posterior knee in the popliteal fossa, while a stabilizing strap was placed over the pelvis to provide stabilization. All participants were able to maintain these positions during testing. Peak force values (kg) were recorded and converted to Newtons (kg*9.81). Torque values (Nm) were calculated by multiplying the force values (N) by the

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Figure 1. Participant positioning and placement of the handheld dynamometer for testing of hip external rotation (top row), abduction (middle row), and extension (bottom row). Sagittal plane hip angle is provided at the top of each individual picture. A positive angle reflects a flexed position; a negative angle reflects an extended position.

limb segment length (m). The femur length was used for hip ABD and EXT calculations and the tibia length was used for hip ER calculations. Torque values were body-size normalized to mass (kg) multiplied by height (m), as recommended by Bazett-Jones et al.29 The three normalized torque values (Nm/kg*m) were averaged for statistical analysis. Body-size normalized torque values were analyzed with a repeated measures ANOVA for hip ABD, ER, and EXT. Following significant effects for sagittal plane hip position, Sidak’s pairwise comparisons

were performed. Statistical significance was set a prior to p<0.05. Effect sizes were calculated using Cohen’s d and the size of the effect was determined using published guidelines: <0.3 = minimal effect, 0.3-0.5 = small effect, 0.5-0.8 medium effect, and >0.8 large effect.32 Statistical analyses were performed in SPSS 24.0 (IBM SPSS Statistics, Version 24, Armonk, NY, USA). RESULTS Table 1 presents means, standard deviations, 95% confidence intervals, p-values, and effect sizes. Peak

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Table 1. Results of normalized peak hip torque measurements across testing positions with varying degrees of hip ﬂexion Sagittal Plane Hip Position 30°

0° Hip Strength Tests Extension Torque (Nm/kg*m)

Mean ± SD

95% CI

0.83 ± 0.30

0.7210.939

External Rotation Torque (Nm/kg*m)

0.22 ± 0.08

Abduction Torque (Nm/kg*m)

0.88 ± 0.17

90°/-5°

Mean ± SD

95% CI

1.14 ± 0.38

1.0021.278

0.1910.249

0.22 ± 0.07

0.8180.942

0.82 ± 0.19

Results

Mean ± SD

95% CI

1.43 ± 0.50

1.2481.612

0.1950.245

0.29 ± 0.13

0.7510.889

0.83 ± 0.24

Effect Size

p-Values Main Effect

<0.001*

0.2430.337

0.027*

0.7430.917

0.085

Pairwise Comparisons

Cohen's d

0 v 30

<0.001*

0.914

0 v 90 30 v 90 0 v 30 0 v 90 30 v 90 0 v 30 0 v -5 30 v -5

<0.001*

1.468

0.002*

0.659

1.000 0.096

0.000 0.654

0.080

0.676

0.066 0.278 0.996

0.336 0.243 0.047

*Significant at p<0.05, SD=standard deviation

torque of ER and EXT were significantly affected by sagittal plane hip position (p<0.05); however, no significant differences were found for peak ABD torque. Post-hoc comparisons revealed differences among testing positions. Peak EXT torque increased as hip flexion increased, with all three positions being significantly different (p<0.05) from one another and having medium to large effect sizes. There was a trend (p<0.10), as well as a medium effect size, toward peak ER torque being greater in 90 of hip flexion than both 0 and 30 . No statistical difference and minimal effect between 0 and 30 for hip ER was found. DISCUSSION The main finding of this study is that hip EXT peak torques are influenced by sagittal plane hip position. Specifically, peak torque values for EXT are greatest at 90 of hip flexion, supporting the hypothesis for EXT. It was hypothesized, based on previous research,19,20 that ER peak torque would be greatest at 90 of hip flexion. This hypothesis was not supported despite finding medium effect sizes between positions. It was hypothesized that ABD peak torque would be greater with 30 of flexion; however, this was not supported since position did not influence peak ABD torque. Given these results, comparison of hip EXT and ER torque values tested in different positions should be performed with caution. The results of this study indicate that peak hip EXT torque increases as hip flexion increases, which is in

agreement with previous research.23,24 Prior studies23,24 have not reported knee position during testing. A flexed knee increases the involvement of the gluteus maximus and reduced the involvement of the hamstring muscles.33,34 With a flexed knee, the results of the current study demonstrate that as hip flexion increases, hip torque increases as well. This is also likely due to an increase in the contribution of nongluteal muscles, specifically the adductors and hamstrings, to hip EXT torque.18 Dostal et al.25 reported that the adductor magnus (5.8 cm), semitendinosis (5.6 cm), biceps femoris (5.4 cm), and semimembranosis (4.6 cm) all have the greater than or equal extensor moment arm lengths in neutral as compared to the gluteus maximus (4.6 cm). As the hip flexes, the gluteus maximus has reduced extensor muscle action and the other hip extensors and adductors contribute to hip extension.25 Gluteus maximus electromyographic analyses have also reported reduced activation as hip flexion increases.35 The torque values reported in this study are the greatest with the hip flexed to 90 , so it is also possible that a flexed hip and flexed knee could place the gluteus maximus and/or hamstrings at an optimal length-tension relationship. This combined with an increased contribution of the adductors could explain the increase hip EXT values with increase flexion. The authors of the current study were unable to find studies investigating the muscle force length-tension relationship in vivo or via musculoskeletal modeling that demonstrate evidence of this. Due to the differences found in hip EXT strength, clinicians should use consistent

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testing positions for patients across sessions and/or between testers. Positions will depend on the goal of strength testing. Testing in a neutral position (0 ) with the knee flexed to 90 may be preferred for targeting the gluteus maximus musculature, while maximum hip EXT muscle group strength may be targeted best by testing in 90 of hip flexion. The results of the current study are in agreement with previously published studies that have reported no differences in ER strength with the hip flexed. Johnson and Hoffman21 reported no differences in hip ER strength among 10 , 40 , and 90 of hip flexion. Bloom and Cornbleet22 also reported no differences between hip ER strength tested in seated (90 hip flexion) and supine (0 hip flexion). In contrast with our results, Lindsay et al.20 reported that peak ER torque was greatest in a seated position (90 hip flexion) compared to supine with the knee flexed or extended. A more recent study by Hoglund et al.19 also reported that the more flexed the participantâ&#x20AC;&#x2122;s hip is during testing, the greater their ER torque value. While some authors19-21 utilized isokinetic dynamometers to test torque, the current study and Bloom and Cornbleet22 used hand-held dynamometry. There are varying results in the literature regarding the effects of hip strength testing positions; however, these results do not seem to be influenced by the devices used. The reason for the variable ER strength differences in 90 compared to 0 of hip flexion have not yet been investigated. Neumann18 utilized data from Delp et al.36 to demonstrate that as the hip is flexed from 0 to 90 , the moment arm of the anterior gluteus maximus and all fibers of the gluteus medius change from causing external rotation to causing internal rotation. Based purely on the moment arm data, we would expect hip ER strength to be reduced in a more flexed position. While hip flexion would reduce the number of muscles or fibers available to produce force, it might improve the length-tension relationship of those fibers that have an ER moment arm, resulting in increased torque values found by some authors.19,20 Future research, possibly involving musculoskeletal modeling, is encouraged to further inform this topic and clarify a mechanism for these findings. This is the first study, to the authorâ&#x20AC;&#x2122;s knowledge, that investigates isometric peak torque production

across different sagittal plane hip positions. It was hypothesized that the hip ABD torque values would be greater with the 30 flexion position due to the addition of tensor fasciae latae contribution, which was not supported. In studies utilizing modeling36 and cadaveric investigations,25 authors have reported that as the hip flexes, the primary abductors (i.e. gluteus medius and minimus) have lower propensity toward abduction due to smaller moment arms and changed lines of action. As the hip flexes, only the tensor fasciae latae and piriformis maintain their abductor line of action from 0 to 90 of hip flexion.25 Authors investigating electromyography of the hip musculature, including the gluteus medius and tensor fasciae latae have reported differences in hip ABD between exercises that involve hip flexion compared to neutral. These authors have reported reduced activation of the gluteus medius during clam exercises that involve 30 -60 of hip flexion.26,27 With reduced gluteus medius activation and increased tensor fasciae latae activation,27 the tensor fasciae latae might contribute more strongly with increased hip flexion. Dostal et al.25 reported that the gluteus medius and minimus maintain an abduction line of action until after 40 of flexion, providing a possible reason that differences were not significant in the current study. Given this discussion (e.g. small moment arm length changes), it seems most appropriate to test hip ABD torque in 0 of hip flexion in order to assess maximum hip ABD strength, however, significant differences were not found in the current study. These findings should be considered within the limitations of this study. As with all studies of maximal muscle contraction, there is an assumption that participants in this study put forth maximal effort. Consistent verbal encouragement was provided in an attempted to reduce the effect of this limitation. This study is limited by the exclusion of electromyography during testing in the various positions, which limits our ability to determine the reasons for differences in strength values. Future investigation of this line of questioning should include electromyographical measurement to evaluate specific muscle activation in different sagittal plane positions. While this study was adequately powered for the comparison of peak ER torque, it may have been underpowered in sample size for the comparison of peak ABD

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torque. Finally, this study is limited by its inclusion of only females and requires further investigation of the influence of sagittal plane hip positon on hip muscle strength in males. Given these limitations, the findings of this study provide good evidence for recommending appropriate positions for hip muscle strength testing in females. CONCLUSION Hip flexion position during hip strength testing has a significant effect on the torque values of the hip musculature. Comparison of strength values between studies that have utilized positions of 0 and 90 hip flexion should be compared with caution, especially for hip EXT. Clinicians and researchers must be careful to conduct and record hip strength testing in the same position within and between sessions. It is also recommended that clinicians choose positions that are most functionally related to the activity of interest or a position that most directly targets the appropriate muscles. Testing all muscles in a neutral position (0 flexion) may allow for efficient testing of all three of these motions during pre-participation examinations or clinical evaluations. REFERENCES 1. Friel K, McLean N, Myers C, et al. Ipsilateral hip abductor weakness after inversion ankle sprain. J Athl Train. 2006;41(1):74-78. 2. Rathleff MS, Rathleff CR, Crossley KM, et al. Is hip strength a risk factor for patellofemoral pain? A systematic review and meta-analysis. Br J Sports Med. 2014;48(14):1088. 3. Pohl MB, Patel C, Wiley JP, et al. Gait biomechanics and hip muscular strength in patients with patellofemoral osteoarthritis. Gait Posture. 2013;37(3):440-444. 4. Verrelst R, Willems TM, De Clercq D, et al. The role of hip abductor and external rotator muscle strength in the development of exertional medial tibial pain: A prospective study. Br J Sports Med. 2014;48(21):1564-1569. 5. Harris-Hayes M, Mueller MJ, Sahrmann SA, et al. Persons with chronic hip joint pain exhibit reduced hip muscle strength. J Orthop Sports Phys Ther. 2014;44(11):890-898. 6. Noehren B, Schmitz A, Hempel R, et al. Assessment of strength, ďŹ&#x201A;exibility, and running mechanics in men with iliotibial band syndrome. J Orthop Sports Phys Ther. 2014;44(3):217-222.

7. Hewett TE, Ford KR, Hoogenboom BJ, et al. Understanding and preventing acl injuries: Current biomechanical and epidemiologic considerations update 2010. N Am J Sports Phys Ther. 2010;5(4):234251. 8. Jaramillo J, Worrell TW, Ingersoll CD. Hip isometric strength following knee surgery. J Orthop Sports Phys Ther. 1994;20(3):160-165. 9. Nadler SF, Malanga GA, DePrince M, et al. The relationship between lower extremity injury, low back pain, and hip muscle strength in male and female collegiate athletes. Clin J Sport Med. 2000;10(2):89-97. 10. Souza RB, Powers CM. Differences in hip kinematics, muscle strength, and muscle activation between subjects with and without patellofemoral pain. J Orthop Sports Phys Ther. 2009;39(1):12-19. 11. Bazett-Jones DM, Cobb SC, Huddleston WE, et al. Effect of patellofemoral pain on strength and mechanics after an exhaustive run. Med Sci Sports Exerc. 2013;45(7):1331-1339. 12. Ferber R, Kendall KD, Farr L. Changes in knee biomechanics after a hip-abductor strengthening protocol for runners with patellofemoral pain syndrome. J Athl Train. 2011;46(2):142-149. 13. Magalhaes E, Fukuda TY, Sacramento SN, et al. A comparison of hip strength between sedentary females with and without patellofemoral pain syndrome. J Orthop Sports Phys Ther. 2010. 14. Herbst KA, Barber Foss KD, Fader L, et al. Hip strength is greater in athletes who subsequently develop patellofemoral pain. Am J Sports Med. 2015;43(11):2747-2752. 15. Boling MC, Padua DA, Alexander Creighton R. Concentric and eccentric torque of the hip musculature in individuals with and without patellofemoral pain. J Athl Train. 2009;44(1):7-13. 16. Boling MC, Padua DA, Marshall SW, et al. A prospective investigation of biomechanical risk factors for patellofemoral pain syndrome: The joint undertaking to monitor and prevent acl injury (jump-acl) cohort. Am J Sports Med. 2009;37(11):21082116. 17. Willy RW, Davis IS. The effect of a hip-strengthening program on mechanics during running and during a single-leg squat. J Orthop Sports Phys Ther. 2011;41(9):625-632. 18. Neumann DA. Kinesiology of the hip: A focus on muscular actions. J Orthop Sports Phys Ther. 2010;40(2):82-94. 19. Hoglund LT, Wong AL, Rickards C. The impact of sagittal plane hip position on isometric force of hip external rotator and internal rotator muscles in

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20.

21.

22.

23.

24. 25. 26.

27.

28.

healthy young adults. Int J Sports Phys Ther. 2014;9(1):58-67. Lindsay DM, Maitland M, Lowe RC, et al. Comparison of isokinetic internal and external hip rotation torques using different testing positions. J Orthop Sports Phys Ther. 1992;16(1):43-50. Johnson S, Hoffman M. Isometric hip-rotator torque production at varying degrees of hip ﬂexion. J Sport Rehabil. 2010;19(1):12-20. Bloom N, Cornbleet SL. Hip rotator strength in healthy young adults measured in hip ﬂexion and extension by using a hand-held dynamometer. PM R. 2014;6(12):1137-1142. Jensen RH, Smith GL, Johnston RC. A technique for obtaining measurements of force generated by the hip muscle. Arch Phys Med Rehabil. 1971;52(5):207215. Clarke HH. Muscular strength and edurance. Englewood Cliffs, NJ: Prentice-Hall; 1966. Dostal WF, Soderberg GL, Andrews JG. Actions of hip muscles. Phys Ther. 1986;66(3):351-361. Distefano LJ, Blackburn JT, Marshall SW, et al. Gluteal muscle activation during common therapeutic exercises. J Orthop Sports Phys Ther. 2009;39(7):532-540. McBeth JM, Earl-Boehm JE, Cobb SC, et al. Hip muscle activity during 3 side-lying hip-strengthening exercises in distance runners. J Athl Train. 2012;47(1):15-23. Thorborg K, Petersen J, Magnusson SP, et al. Clinical assessment of hip strength using a hand-held dynamometer is reliable. Scand J Med Sci Sports. 2010;20(3):493-501.

29. Bazett-Jones DM, Cobb SC, Joshi MN, et al. Normalizing hip muscle strength: Establishing body-size-independent measurements. Arch Phys Med Rehabil. 2011;92(1):76-82. 30. Bolgla LA, Malone TR, Umberger BR, et al. Hip strength and hip and knee kinematics during stair descent in females with and without patellofemoral pain syndrome. J Orthop Sports Phys Ther. 2008;38(1):12-18. 31. Widler KS, Glatthorn JF, Bizzini M, et al. Assessment of hip abductor muscle strength. A validity and reliability study. J Bone Joint Surg Am. 2009;91(11):2666-2672. 32. Cohen J. Statistical power analysis for the behavioral sciences. 2nd ed1988. 33. Sakamoto ACL, Teixeira-Samela LF, Rodrigues de Paula F, et al. Gluteus maximus and semitendinosus activation during active prone hip extension exercises. Rev Bras Fisioter. 2009;14(4):335-342. 34. Jeon IC, Kwon OY, Weon JH, et al. Comparison of hip and back muscle activity and pelvic compensation in healthy subjects during three different prone table hip extension exercises. J Sport Rehabil. 2016. 35. Worrell TW, Karst G, Adamczyk D, et al. Inﬂuence of joint position on electromyographic and torque generation during maximal voluntary isometric contractions of the hamstrings and gluteus maximus muscles. J Orthop Sports Phys Ther. 2001;31(12):730740. 36. Delp SL, Hess WE, Hungerford DS, et al. Variation of rotation moment arms with hip ﬂexion. J Biomech. 1999;32(5):493-501.

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IJSPT

ORIGINAL RESEARCH

BUILDING A BETTER GLUTEAL BRIDGE: ELECTROMYOGRAPHIC ANALYSIS OF HIP MUSCLE ACTIVITY DURING MODIFIED SINGLE-LEG BRIDGES B.J. Lehecka, DPT1 Michael Edwards1 Ryan Haverkamp1 Lani Martin1 Kambry Porter1 Kailey Thach1 Richard J. Sack2 Nils A. Hakansson, PhD3

ABSTRACT Background: Gluteal strength plays a role in injury prevention, normal gait patterns, eliminating pain, and enhancing athletic performance. Research shows high gluteal muscle activity during a single-leg bridge compared to other gluteal strengthening exercises; however, prior studies have primarily measured muscle activity with the active lower extremity starting in 90° of knee flexion with an extended contralateral knee. This standard position has caused reports of hamstring cramping, which may impede optimal gluteal strengthening. Hypothesis/Purpose: The purpose of this study was to determine which modified position for the single-leg bridge is best for preferentially activating the gluteus maximus and medius. Study Design: Cross-Sectional Methods: Twenty-eight healthy males and females aged 18-30 years were tested in five different, randomized single-leg bridge positions. Electromyography (EMG) electrodes were placed on subjects’ gluteus maximus, gluteus medius, rectus femoris, and biceps femoris of their bridge leg (i.e., dominant or kicking leg), as well as the rectus femoris of their contralateral leg. Subjects performed a maximal voluntary isometric contraction (MVIC) for each tested muscle prior to performing five different bridge positions in randomized order. All bridge EMG data were normalized to the corresponding muscle MVIC data. Results: A modified bridge position with the knee of the bridge leg flexed to 135° versus the traditional 90° of knee flexion demonstrated preferential activation of the gluteus maximus and gluteus medius compared to the traditional single-leg bridge. Hamstring activation significantly decreased (p < 0.05) when the dominant knee was flexed to 135° (23.49% MVIC) versus the traditional 90° (75.34% MVIC), while gluteal activation remained similarly high (51.01% and 57.81% MVIC in the traditional position, versus 47.35% and 57.23% MVIC in the modified position for the gluteus maximus and medius, respectively). Conclusion: Modifying the traditional single-leg bridge by flexing the active knee to 135° instead of 90° minimizes hamstring activity while maintaining high levels of gluteal activation, effectively building a bridge better suited for preferential gluteal activation. Level of Evidence: 3 Key words: Gluteus maximus, gluteus medius, muscle recruitment, rehabilitation exercise

Department of Physical Therapy, Wichita State University, Wichita, KS, USA 2 Department of Electrical Engineering and Computer Science, Wichita State University, Wichita, KS, USA 3 Department of Biomedical Engineering, Wichita State University, Wichita, KS, USA Funding: Funding was provided by Kansas Partners in Progress (KPIP) Conflict of Interest: There are no potential conflicts of interests, including financial arrangements, organizational affiliations, or other relationships that may constitute a conflict of interest regarding the submitted work.

CORRESPONDING AUTHOR BJ Lehecka, DPT 1845 Fairmount St. Wichita, KS 67260 Phone: 316-978-6156 E-mail: bryan.lehecka@wichita.edu

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INTRODUCTION Gluteal muscle strength and endurance play a significant role in injury prevention, normalizing gait patterns and posture, eliminating pain, and enhancing athletic performance.1-5 Weakness of the hip abductors and external rotators appears to be a risk factor for injury in collegiate and track athletes.5 Gluteus medius strengthening has been shown to improve functional recovery and pain reduction in patients following knee meniscus surgery.1 Gluteus medius endurance and active hip abduction tests are predictive of individuals at risk for low back pain during prolonged standing.2,3 Additionally, gluteus maximus strengthening has resulted in a decrease in low back pain and disability.4 The strength of these two gluteal muscles, which comprise about 33% of the hip musculature,6 is essential for athletic, non-athletic, and post-surgical populations. Gluteal muscle activation during common therapeutic exercises has been the focus of numerous studies.7-11 The single-leg bridge can yield sufficient gluteus maximus and gluteus medius muscle activity for strengthening without external loading and thereby provide individuals a means to safely and conveniently increase hip joint stability. Prior electromyographic (EMG) analysis of hip muscles during the single-leg bridge demonstrated 40% activation when normalized to the maximum voluntary isometric contraction (MVIC) for the gluteus maximus, and 47% MVIC for the gluteus medius.11 The single-leg bridge produced the secondhighest activation of gluteal muscles among the nine rehabilitation exercises examined. In the same study, the single-leg bridge produced the highest level of hamstring activation (40% MVIC).11 A common clinical finding during the single-leg bridge is hamstring cramping, perhaps due to the combination of gluteal muscle weakness and high hamstring activation. In clinical practice, hamstring cramping during the single-leg bridge often prevents sufficient repetitions of the exercise to be performed, which may diminish the contribution to gluteal strengthening. In one EMG study of gluteal exercises, there were multiple subject reports of hamstring cramping during the single-leg bridge on both stable and unstable surfaces.9 To the authors’ knowledge, EMG activity of hip muscles during the single-leg bridge has primarily been

studied in the traditional starting position characterized by 90° of knee flexion on the stance leg, with the ipsilateral foot flat and the contralateral knee fully extended. Based on clinical experience and subjective reports from patients and athletes, the authors hypothesized that alterations to the singleleg bridge exercise could result in more efficacious gluteal muscle outcomes. Therefore, the purpose of this study was to determine which modified position for the single-leg bridge is best for preferentially activating the gluteus maximus and medius. METHODS Twenty-eight subjects (16 females, 12 males) were recruited for this cross-sectional study from a sample of convenience. The average age was 23.43 ± 2.28 years. The average height and weight were 1.73 ± 0.11 meters and 72.57 ± 13.93 kilograms, respectively. Healthy subjects between 18 and 30 years of age able to perform one hour of low-intensity exercise completed testing in a biomechanics laboratory at the local university. Subjects completed an informed consent form and health questionnaire prior to testing and were excluded if they had any of the following: current pain or pain with exercise in their low back or lower extremity; numbness or tingling in their low back or lower extremity; back or lower extremity surgery within the past two years; pacemaker; or current pregnancy. Subjects were familiarized with the testing procedures, including the familiarization trial, electrode placement, stationary bicycle warm-up, MVIC testing, and single-leg bridge variations. Subjects were asked to remove shoes and perform the bridging exercises in socks. Subjects underwent a familiarization trial for each of the five variations of bridges, and performed two repetitions of each variation prior to formal testing. Surface EMG electrodes were used to record muscle activity as the subjects performed the single–leg bridge exercises. Prior to electrode placement, the skin was shaved and abraded with alcohol wipes. Surface electrodes were placed by two researchers. One used a standard tape measure to identify landmarks for electrode placement, while the other placed the electrodes. Surface EMG data were collected at 3000 Hz using a Noraxon TeleMyo 2400T

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GT (Noraxon, Scorrsdale, AZ). Bi-polar electrodes were placed on the quadriceps femoris, biceps femoris, gluteus maximus, and gluteus medius of the bridge leg, i.e., the dominant leg as determined by which leg would be used to kick a ball. One electrode was placed on the contralateral quadriceps femoris. This was the only electrode placed contralaterally because the only modification made to that limb (simultaneous hip flexion to a vertical position and passive knee flexion) was intended to reduce activity in this quadriceps muscle. Specific electrode placement is described in Table 1 and was based on similar studies and standard practice.11-13 Subjects pedaled a stationary bicycle at 60 rpm with a work rate of 60 W for five minutes as a warm-up. Following the warm-up, electrodes were placed in the aforementioned positions and secured with paper tape to ensure adherence to the skin. Muscle MVIC testing was then performed in the following order: ipsilateral gluteus medius, ipsilateral gluteus maximus, ipsilateral biceps femoris, ipsilateral rectus femoris, and contralateral rectus femoris. A strap attached to an immobile object was used to standardize resistance during MVIC testing. Descriptions of positions used for MVIC testing are displayed in Table 2. Subjects completed three MVIC trials for

each muscle. Subjects were asked to complete each MVIC for seven seconds, with 30 seconds of rest between each trial.11 Each subject performed five variations of the single-leg bridge. The positions of these five variations (Positions A-E) are described in Table 3 and Figures 1-5. Position order was randomized, and the data collector was blind to the position being performed. Subjects were blind to the EMG activity of their hip muscles. Subjects performed eight trials of each bridge variation to the beat of a metronome set at 60 beats per minute, extending their hip to neutral during each trial. Recorded data from trials three through seven were post-processed and included in the data analysis. All EMG data were rectified and filtered using a 15 Hz high-pass and 500 Hz low-pass fourth-order Butterworth digital filter. The filtered EMG data were smoothed with a 50-millisecond moving average. Data plots were then inspected twice by a set of five researchers to remove movement artifacts. For each position and each participant, five bridges were recorded and five %MVICs were calculated. MVIC values for each muscle were identified as the maximum value of a 50-millisecond moving

Table 1. Electrode Placements

Table 2. MVIC Testing Positions

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Table 3. EMG Activity of Hip Muscles in Different Single-Leg Bridge Positions (Mean ± SD of %MVIC)

Figure 1. Position A. Subjects started with their dominant knee ﬂexed to 90° and foot ﬂat. The contralateral knee was extended and its hip remained in neutral. Arms were folded across the chest.

Figure 2. Position B. Subjects started with their dominant knee ﬂexed to 135° and foot ﬂat. The contralateral knee was extended and its hip remained in neutral. Arms were folded across the chest.

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Figure 3. Position C. Subjects started with their dominant knee flexed to 90° and foot flat. The contralateral knee was relaxed in flexion with the femur held vertical. Arms were folded across the chest.

Figure 5. Position E. Subjects started with their dominant knee flexed to 135° and ankle in full dorsiflexion. The contralateral knee was relaxed in flexion with the femur held vertical. Arms were folded across the chest.

RESULTS Data from twenty-six subjects were included. Data from two subjects were excluded due to faulty data from the EMG leads for the gluteal muscles. Means and standard deviations of EMG activity expressed as the %MVIC of each analyzed muscle in each of the five bridge positions are presented in Table 3. Significant differences in muscle activation between the modified bridge positions and traditional bridge position (position A) are also noted in Table 3.

Figure 4. Position D. Subjects started with their dominant knee flexed to 90° and ankle in full dorsiflexion. The contralateral knee was relaxed in flexion with the femur held vertical. Arms were folded across the chest.

average within the three corresponding seven–second MVIC trials. The mean and standard deviation values of these %MVICs were used in final analysis. All data processing was performed using custom written code (Matlab, The Mathworks, Natick, MA). Repeated measures analyses of variance (ANOVAs) (α = 0.05) with Bonferroni corrections were performed for each muscle tested in the five bridge variations. SPSS v23.0 (SPSS Inc, Chicaco, IL) was used for data analysis.

Hamstring activity (i.e., percent MVIC) was minimized in positions B (23.49%) and E (20.84%), in which the dominant knee was flexed to 135° instead of 90°. These positions appeared to preferentially activate both gluteal muscles, as gluteus maximus and gluteus medius activation surpassed and essentially doubled biceps femoris activation. Of the two positions, position B displayed higher %MVIC for both gluteal muscles (47.35% and 57.23% versus 40.38% and 41.63% for the gluteus maximus and gluteus medius, respectively). Gluteus maximus and gluteus medius percent MVIC in position E (40.38% and 41.63%, respectively) was significantly (p < 0.05) less than position A (51.02% and 57.81%, respectively), the traditional single-leg bridge position. Gluteus maximus and gluteus medius activation in position B (47.35% and 57.23%, respectively), however, did not significantly change compared to position A.

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DISCUSSION The primary objective of this study was to examine hip muscle activity during five variations of the single-leg bridge and determine which variation preferentially activated the gluteal muscles. The positions and procedures of this study were similar to those used by Boren et al.9 and Ekstrom et al.11 Boren et al. found the traditional single-leg bridge elicited 54% MVIC of the gluteus maximus and 54% MVIC of the gluteus medius.9 Ekstrom et al. found the traditional single-leg bridge elicited 40% MVIC of the gluteus maximus and 47% MVIC of the gluteus medius.11 These data are similar to the results of the current study which showed 51.01% MVIC for the gluteus maximus and 57.81% MVIC for the gluteus medius in the traditional single-leg bridge position (position A). A difference was seen between EMG activity of the bridge leg biceps femoris during the traditional single-leg bridge (position A) in this study (75.34% MVIC) and that found by Ekstrom et al. (40% MVIC).11 As the MVIC of the biceps femoris was achieved in a similar manner, the difference in activation may be due to the variation in upper extremity placement. Subjects in the study by Ekstrom et al. placed their upper extremities flat on the testing surface to their side.11 Subjects in the current study were instructed to fold their arms across the chest to ensure they did not aid lower extremity muscle activity during the exercise. This position difference may explain the difference in biceps femoris activity between studies. Placement of the upper extremities by the side would have permitted the subjects to use the arms to apply a downward force on the testing surface. Such a force would have been manifested as an extensor moment about the shoulder joints that would have reduced the weight force supported by the bridge leg and thereby reduced the biceps femoris muscle activity required to maintain the bridge position. Folding the arms across the chest prevented the subjects from generating an extensor moment about the shoulder joints that contributed to maintaining the bridge position. Significant decreases in biceps femoris activity occurred in this study when the knee was flexed to 135° versus 90°. Biceps femoris activity decreased from 75.34% in the traditional position (position A) to 23.49% in position B. This significant decrease in

biceps femoris activity following increased knee flexion may be due to the lower leg being aligned parallel with the ground reaction force vector at the foot. In this case, there is a reduced knee extensor moment and a reduced need for the hamstrings to actively maintain the knee flexion angle during the bridge. Reduced hamstring activity would decrease the incidence of hamstring cramping, as one theory of exercise-associated muscle cramps is that they occur due to muscle overload and neuromuscular fatigue.14 Muscle activation during the bridge with a starting position of 135° knee flexion is unable to be compared to other studies as it appears to be a novel starting position. A study by Youdas et al. describes EMG activity of a single-leg bridge that requires a subject to achieve a 90° flexion angle of the knee at the height of the bridge, presumably starting the exercise with a higher degree of knee flexion.15 However, the starting knee angle of that single-leg bridge was not measured or described. Subjects in that study were also allowed to place their hands on the floor by the sides, potentially assisting the activity. Significant changes in hip muscle activation occurred with different ankle positioning. Dorsiflexing the ankle of the dominant leg appeared to significantly decrease biceps femoris activity in the current study. Dorsiflexion of the ankle was the only difference between positions C and D, and biceps femoris activity significantly decreased from 69.18% MVIC (position C) to 58.71% MVIC (position D) when the ankle was dorsiflexed. With ankle dorsiflexion, the gastrocnemius was lengthened and thereby no longer actively insufficient to generate a knee flexion moment. Gastrocnemius activity would reduce the need for the hamstrings to be active to prevent knee extension. Additionally, in a study by Chon et al., ankle dorsiflexion was shown to enhance transversus abdominus activity during an abdominal draw-in maneuver (i.e., hook-lying position).16 Thus, ankle dorsiflexion may have caused a decrease in biceps femoris activity during the bridge by activating alternate core musculature on the anterior side of the body, in turn lessening the demand on the hamstring muscles posteriorly. In position E, a significant decrease in gluteal activity compared to position A was observed when dorsiflexion, increased knee flexion to 135°, and contralateral hip flexion were all added to the exercise.

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In the current study, exercise order was randomized and both researchers and subjects were blinded when possible, yet limitations remain. There is potential that subjects did not generate a true MVIC of each muscle tested due to lack of effort or suboptimal positioning. Muscle length during MVIC testing may also be a factor. Attempts were made to minimize these potential limitations by standardizing instructions to subjects, standardizing positioning methods, and using methods for the traditional single-leg bridge and MVIC testing similar to previous studies. Hip muscle activity during single-leg bridge positions was not studied in subjects with pathology, so generalizing results to an injured population should be performed with caution. Future research should examine the effects of strength training using the modified single-leg bridge with preferential gluteal activation (position B) and its effects on pathology and performance. CONCLUSION The modified single-leg bridge position with 135° of knee flexion (position B) displayed preferential activation of the gluteus maximus and gluteus medius over the biceps femoris. This position maintained gluteal activity while significantly decreasing biceps femoris activity compared to the traditional singleleg bridge position. This modified single-leg bridge, potentially called “the gluteal bridge,” using 135° of knee flexion on the dominant side may allow more optimal training of the gluteal muscles than the traditional position since the biceps femoris is less likely to be the limiting muscle of the exercise. REFERENCES 1. Kim EK. The effect of gluteus medius strengthening on the knee joint function score and pain in meniscal surgery patients. J Phys Ther Sci. 2016;28(10):2751-2753. 2. Nelson-Wong E, Flynn T, Callaghan JP. Development of active hip abduction as a screening test for identifying occupational low back pain. J Orthop Sports Phys Ther. 2009;39(9):649-57. 3. Marshall PW, Patel H, Callaghan JP. Gluteus medius strength, endurance, and co-activation in the development of low back pain during prolonged standing. Hum Mov Sci. 2011;30(1):63-73. 4. Jeong U-C, Sim J-H, Kim C-Y, et al. The effects of gluteus muscle strengthening exercise and lumbar

stabilization exercise on lumbar muscle strength and balance in chronic low back pain patients. J Phys Ther Sci. 2015;27(12):3813-3816. 5. Leetun DT, Ireland ML, Willson JD, et al. Core stability measures as risk factors for lower extremity injury in athletes. Med Sci Sports Exerc. 2004;36(6):926-34. 6. Ito J, Moriyama H, Inokuchi S, et al. Human lower limb muscles: an evaluation of weight and fiber size. Okajimas Folia Anat Jpn. 2003;80(2-3):47-55. 7. Distefano LJ, Blackburn JT, Marshall SW, et al. Gluteal muscle activation during common therapeutic exercises. J Orthop Sports Phys Ther. 2009;39(7):532-40. 8. Reiman MP, Bolgla LA, Loudon JK. A literature review of studies evaluating gluteus maximus and gluteus medius activation during rehabilitation exercises. Physiother Theory Pract. 2012;28(4):257-68. 9. Boren K, Conrey C, Le Coguic J, et al. Electromyographic analysis of gluteus medius and gluteus maximus during rehabilitation exercises. Int J Sports Phys Ther. 2011;6(3):206-23. 10. Selkowitz DM, Beneck GJ, Powers CM. Which exercises target the gluteal muscles while minimizing activation of the tensor fascia lata? Electromyographic assessment using fine-wire electrodes. J Orthop Sports Phys Ther. 2013;43(2):54-64. 11. Ekstrom RA, Donatelli RA, Carp KC. Electromyographic analysis of core trunk, hip, and thigh muscles during 9 rehabilitation exercises. J Orthop Sports Phys Ther. 2007;37(12):754-62. 12. Cram JR, Kasman G. Introduction to Surface Electromyography. Gaithersburg, MD: Aspen Publishers, Inc; 1997. 13. Finni T, Cheng S. Variability in lateral positioning of surface EMG electrodes. J Appl Biomech. 2009;25:396400. 14. Miller KC, Stone MS, Huxel KC, et al. ExerciseAssociated Muscle Cramps: Causes, Treatment, and Prevention. Sports Health. 2010;2(4):279-283. 15. Youdas JW, Hartman JP, Murphy BA, et al. Magnitudes of muscle activation of spine stabilizers, gluteals, and hamstrings during supine bridge to neutral position. Physiother Theory Pract. 2015;31(6):418-27. 16. Chon SC, Chang KY, You JS. Effect of the abdominal draw-in manoeuvre in combination with ankle dorsiflexion in strengthening the transverse abdominal muscle in healthy young adults: a preliminary, randomised, controlled study. Physiotherapy. 2010;96(2):130-6.

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IJSPT

ORIGINAL RESEARCH

EFFECTS OF A BAND LOOP ON LOWER EXTREMITY MUSCLE ACTIVITY AND KINEMATICS DURING THE BARBELL SQUAT Ryan C.A. Foley1 Brittany D. Bulbrook1 Duane C. Button2 Michael W.R. Holmes1,3

ABSTRACT Background: Medial knee collapse can signal an underlying movement issue that, if uncorrected, can lead to a variety of knee injuries. Placing a band around the distal thigh may act as a proprioceptive aid to minimize medial collapse of the knee during squats; however, little is known about EMG and biomechanics in trained and untrained individuals during the squat with an elastic band added. Hypothesis/Purpose: To investigate the effects of the TheraBand® Band Loop on kinematics and muscle activity of the lower extremity during a standard barbell back squat at different intensities in both trained and untrained individuals. Study Design: Cross-sectional, repeated measures. Methods: Sixteen healthy, male, university aged-participants were split into two groups of eight, consisting of a trained and untrained group. Participants performed both a 3-repetition maximum (3-RM) and a bodyweight load squat for repetitions to failure. Lower extremity kinematics and surface electromyography of four muscles were measured bilaterally over two sessions, an unaided squat and a band session (band loop placed around distal thighs). Medial knee collapse, measured as a knee width index, and maximum muscle activity were calculated. Results: During the 3-RM, squat weight was unaffected by band loop intervention (p=0.486) and the trained group lifted more weight than the untrained group (p<0.007). The trained group had a greater squat depth for both squat conditions, regardless of the band (p=0.0043). Knee width index was not affected by the band during the eccentric phase of bodyweight squats in the trained (band: 0.76 ± 0.08, no band: 0.73 ± 0.08) or untrained group (band: 0.77 ± 0.70, no band: 0.75 ± 0.13) (p=0.670). During the concentric phase, knee width index was significantly lower for 3-RM squats, regardless of group. Conclusion: Despite minimal changes in kinematics for the untrained group, increased muscle activity with the band loop may suggest that a training aid may, over time, lead to an increase in barbell squat strength by increasing activation of agonist muscles more than traditional, un-banded squats. Greater maximal muscle activity in most muscles during band loop sessions may provide enhanced knee stability via increased activation of stabilizing muscles. Level of Evidence: 3 Key words: Elastic resistance, electromyography, knee valgus, kinematics, squat, TheraBand

Faculty of Health Sciences, University of Ontario Institute of Technology, Oshawa, Ontario, Canada 2 School of Human Kinetics and Recreation, Memorial University of Newfoundland, St. John’s, Newfoundland, Canada 3 Neuromuscular Mechanics and Ergonomics Laboratory, Department of Kinesiology, Brock University, St. Catharines, Canada

CORRESPONDING AUTHOR Michael W.R. Holmes, PhD Assistant Professor Brock University Department of Kinesiology Niagara Region, 1812 Sir Isaac Brock Way St. Catharines, ON L2S 3A1, brocku.ca Phone: 905 688 5550 x Fax: 905 984 4851 E-mail: michael.holmes@brocku.ca

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INTRODUCTION During a barbell squat, there is a tendency for novice or untrained users to have a medial collapse of the knees under load, especially during the concentric phase of movement. Medial knee collapse, or knee valgus, occurs when there is excessive medial movement of the knee in the frontal plane.1 It can be the result of valgus force, in the frontal plane, that ultimately results in the knee joint center moving towards the midline of the body.2,3 This knee movement will generally be away from alignment with the vertical ground reaction force (GRF) vector,4,7 resulting in hip adduction and internal rotation. Medial collapse is often identified in resistance training activities such as the barbell squat, during explosive jumping,8,9 landing,4,9 and cutting8 tasks. Medial knee collapse can signal an underlying movement issue that, if uncorrected, can lead to a variety of knee injuries, including patellofemoral syndrome,11,3 anterior cruciate ligament tears12 and degeneration of the meniscus due to altered compressive forces on the femur and the patella.1 Coordinated muscle action and movement patterns are essential for optimizing barbell squat efficiency and reducing injury risk. Increasing the load on a barbell squat when medial knee collapse is already evident will likely increase medial knee loading and exacerbate the risk of injury. Previous authors have suggested that hip muscle activity, particularly gluteus maximus, is directly linked to valgus knee angles and internal hip rotation.1,13 High activity in all the gluteal muscles prior to concentric movements produces a robust (and stable) linked system through the hip and knee, resisting valgus knee collapse, and reducing injury risk.13 Quadriceps/hamstring strength ratios may also contribute to valgus knee moments, as both weak hamstrings14 and reduced muscle cocontraction15 can influence valgus knee moments and internal tibial rotation. Lastly, distal movement dysfunction in the kinetic chain, such as decreased ankle dorsiflexion can both increase valgus knee loads and decrease quadriceps activation.16 These mechanisms of valgus knee loading suggest that a detailed biomechanical analysis is required to fully determine the effect of elastic resistance on squat performance. Training aids include a category of fitness equipment designed to coach proper movement patterns through various types of biofeedback or proprio-

ception. These types of tools are popular in sports and rehabilitation that demand complex movement patterns for repeated success in sport and work. Strength coaches and fitness professionals use training aids to more quickly and efficiently correct dysfunctional movements in athletes/patients. Previous research has investigated the use of an elastic loop, or band, around the distal thigh during body weight squats.7 Gooyers et al.12 concluded that the band was unsuccessful at correcting medial collapse or valgus moments; however, the authors acknowledged that results may be task and/or context specific. It is worth noting that the participant population in that study12 was recreationally active students, likely familiar with, but not well trained in proper squat technique and mechanics. Barbell squats provide a similar movement as the body weight squat investigated by Gooyers et al.,12 however, much greater load bearing is required. It is possible that underlying mechanical deficiencies could be exacerbated when the lower extremity is placed under greater load. If a loaded barbell squat results in greater medial knee collapse, this scenario may be more conducive to determining if using a resistance band can be parlayed into a biomechanical benefit. If a band placed around the distal thighs can affect squat mechanics, then it is likely that muscle activity is also affected. Gooyers and colleagues did not measure lower extremity muscle activity during their investigation. Therefore, the purpose of this study was to investigate the effects of the TheraBandÂŽ Band Loop on kinematics and muscle activity of the lower extremity during a standard barbell back squat at different intensities in both trained and untrained individuals. It was hypothesized that the trained groupsâ&#x20AC;&#x2122; kinematics (medial knee collapse) would be unaffected by the band, while the untrained group would experience medial knee alignment changes when using the band. Additionally, it was hypothesized that external hip rotator muscle activity would be greater for the trained participants during the band condition. METHODS Participants Sixteen male, university aged-participants were split into two groups of eight: Trained (25.4 Âą 4.4 years,

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179.83 ± 8.81 cm, 88.36 ± 12.52 kg) and untrained (22.8 ± 1.6 years, 180.81 ± 6.04 cm, 76.91 ± 9.29 kg). The trained group was defined as regularly participating in barbell back squat training for the last year. The untrained group consisted of individuals with no barbell back squatting experience, but who were able to squat a bodyweight load. All participants completed a PAR-Q+ to determine physical readiness for exercise as well as a custom exercise history questionnaire and had no previous history of musculoskeletal pain or injury in the past 12 months. Participants provided informed consent and the study was approved by the University of Ontario Institute of Technology Research Ethics Board (REB#:14-057) and in accordance with the declaration of Helsinki. Protocol This study was designed as a mixed-model repeated measures, with both groups (trained and untrained) performing two squat conditions during two separate sessions (Figure 1). The first session included the control or no band day, where squats were performed with no instruction given to participants. After a minimum of 48 hours rest, but no more than four weeks, participants repeated the protocol with a red, medium-resistance Theraband® Band Loop

(The Hygenic Corporation, OH, USA) placed around the distal thigh, just proximal to the lateral epicondyle of the femur (Figure 2). Trained and untrained participants were given the same coaching on the use of the band; an explanation that the intent is to “keep the band tight throughout the entire squat” but limited verbal cues were given during the actual protocol. Aside from the Theraband® Band Loop intervention, both sessions followed an identical protocol, using the medium resistance band that requires 4.5lbs of pull to stretch a 12-inch band to 24-inches. Data collection began with muscle specific isometric maximal voluntary contractions (MVCs) for each of the eight muscles recorded. Two MVCs were collected for each muscle, with rest given between trials. Participants then performed a warm up consisting of five minutes of cycling at 75 watts. Participants first performed a 3-repetition maximum (3-RM) protocol, starting with a weight approximately half of their estimated 1-repetition maximum, and taking no less than three sets, and no more than five sets to attain a true 3RM. A ‘true’ 3RM was noted as achieved when the participant attempted a further set at the smallest weight increment possible (10 lbs), ending in failure

Figure 1. Timeline of protocol. Each participant (trained or untrained) participated in two data collection sessions. The ﬁrst session consisted of background information, set-up, calibration, normalization, and a 3RM and BW squat protocol. No earlier than 48 hours later the participants participated in an identical bout of data collection, using a band placed around the proximal knees. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 552

Figure 2. Left: EMG and rigid body placement during the no band session. Middle: Band placement during the band session. Right: Anatomical reconstruction and Visual3D model of participant performing a squat.

(referred to as 3RM condition). Rest periods between sets were participant determined and regulated to two to four minute breaks. After a mandatory 10-minute rest period, participants completed a second squat condition, this time with a bodyweight load for maximum repetitions to failure (referred to as BW condition). Heavy encouragement was given to ensure maximum effort and repetitions for each participant. Termination of the BW condition was determined by failure to squat the load, requiring assistance, or a significant deviation in tempo or form with interrepetition pauses exceeding three seconds. Foot position was the participant’s natural squatting stance and was not standardized in order to avoid altering incoming mechanics and finally, to minimize injury risk when performing 3RM repetitions. Muscle Activity Muscle activity was recorded from four muscles bilaterally: gluteus maximus (GMa), gluteus medius (GMe), vastus lateralis (VL) and biceps femoris (BF). Ag-AgCl disposable electrodes (MediTrace 130, Kendall, Mansfield, MA, USA) were placed over each muscle belly in-line with muscle fiber orientation, according to previous work.17 Prior to electrode placement, muscle specific locations were shaved, skin was abraded using NuPrep abrasive gel (Weaver & Company Inc., CO, USA) and cleaned with an isopropyl alcohol. Electromyography (EMG) signals were differentially amplified, band pass filtered (CMRR > 115 dB at 60Hz; input impedance ~10GΩ; 10-1000

Hz; AMT-8, Bortec Biomedical Ltd., Calgary, AB, Canada) and sampled at 2000 Hz. All MVCs included the participant exerting a 3-second maximal effort contraction. Two MVCs were performed for each muscle. Vastus lateralis MVCs were conducted with the participant seated and the leg positioned at 90˚ and secured using an ankle cuff and non-deforming steel cable for a maximal isometric knee extension effort. Biceps femoris used the same procedure but the knee flexion contraction was resisted at the ankle by the experimenter. Gluteus maximus MVCs were conducted with the participant prone on a padded exercise mat. The experimenter resisted the thigh segment while the participant maximally extended their hip, while maintaining knee flexion of 90˚. Gluteus medius MVCs were conducted with the participant lying on their side with both knees flexed to 90˚. The hip of the participants’ top leg was maximally abducted against experimenter resistance. Kinematics and Ground Reaction Forces 3D kinematics were collected using three 3D Investigator Active Motion Capture Systems (Northern Digital Inc., Waterloo, ON, Canada). Custom rigid bodies, consisting of at least three non-collinear markers were placed on the participant’s foot, shank and thigh bilaterally as well as the pelvis and thorax using double sided carpet tape (3M, London, ON, Canada) and Hypafix® (BSN Medical Inc., Hamburg, Germany). Anatomical landmarks were digitized on each participant, assuming a fixed spatial relationship with

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the rigid body affixed to each segment. Kinematics were sampled at 50 Hz and synchronized with EMG data. The global coordinate system was determined as X, medial-lateral, Y, anterior-posterior and Z, superior-inferior. Data Analysis EMG was full wave rectified and Butterworth low pass filtered (3Hz cut-off, dual pass, 2nd order). Peak activity was determined from each muscle specific MVC and muscle activity during each condition was normalized as a percentage of maximal voluntary contraction (%MVC). Kinematic data was used to determine the start, bottom (maximum knee flexion) and end of each squat repetition such that mean and maximum muscle activity could be determined for the concentric and eccentric phases (MatLab 2015b, Mathworks Inc., Natick, MA, USA). The concentric and eccentric phases were determined as frame numbers that marked peak knee flexion and peak knee extension. Additionally, knee angle end points were corroborated using absolute coordinate system data from the thigh markers. Affirming that peak knee-flexion angle occurred at the frame that corresponded to the lowest vertical marker distance. EMG was synchronized with kinematics, so the kinematic frame numbers representing each phase were corrected for sampling rate differences and used for EMG analysis. Kinematic data were processed using Visual3D (C-Motion, Germantown, MD, USA). Local anatomical frames of reference were created for each segment and used in the kinematic and kinetic calculations. Raw kinematic data were low pass Butterworth filtered at 6Hz. Knee joint angles were calculated as the thigh relative to the shank, using an XYZ rotation sequence. Maximum and minimum knee flexion angles were used to corroborate the start of the eccentric and concentric phases for each repetition. From these angles, the distal joint coordinates (XYZ) were calculated for each segment. Knee Width Index (KWI) was calculated from the threedimensional position data from each segment as the ratio of the distance between the right and left distal thigh, and ankle (distal shank points).4,7,18 This analysis was replicated for each repetition of the 3RM and body weight (BW) squats for both the band and no band conditions.

Statistical Analyses An a priori power analysis was performed using G-Power software indicating a need for 16 participants, achieving an actual power of 0.97.5,6 Statistical analyses were conducted using Statistica® (Dell Software Inc., Nashua, NH, USA). A mixed model, General Linear Measures ANOVA was used to determine effect of training (Trained or Untrained), rep/load (3RM or BW), and condition (Band or No Band). All main effects of group or interaction were tested for significance using an alpha of p<0.05 determined a priori. Planned comparisons between conditions for KWI, average and maximum muscle activity, were conducted on Band-Group interaction. Mixed-model ANOVAs (group x squat type x band condition) were performed for knee angle (maximum knee flexion), KWI (bottom) and EMG for each muscle bilaterally. All data are presented as mean ± SD. RESULTS Squat Kinematics There was no effect of band on weight lifted for 3RM intensity (trained, band: 132.7 ± 21.2 kg, no band: 130.3 ± 19.5 kg, p=0.758; untrained, band: 104.5 ± 10.8 kg, no band: 104.0 ± 9.7 kg, p=0.486). The trained group lifted significantly more weight than the untrained group for both the band (p=0.005) and no band (p=0.007) conditions. There were no significant differences in the number of repetitions performed by the trained participants during the BW band and no band conditions (band: 17.7 ± 9.1 repetitions; no band: 18.1 ± 10.0 repetitions, p=0.935). There were also no significant differences in the number of repetitions performed by the untrained participants during the BW band and no band conditions (band: 14.9 ± 9.5 repetitions; no band: 16.1 ± 6.5 repetitions, p=0.762). The trained group had a significantly greater squat depth (knee flexion angle) than the untrained group (trained: 100.94 ± 13.63°; untrained: 89.40 ± 15.03°; p=0.0043) for both 3RM and BW conditions, regardless of the band. Participant and squat demographics can be found in Table 1. Knee Width Index KWI was not affected by the band during the eccentric phase of 3RM squats in either the trained (band: 0.69 ± 0.12, no band: 0.71 ± 0.14) or untrained group

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Table 1. Summary of squat outcomes

Squat Condition 3RM BW

Training Status Trained Untrained Trained Untrained

No band Weight (kg) 133.4 ± 23.4 102.2 ± 4.4 91.2 ± 12.6 76 ± 8.9

Reps 3.0 3.0 15.7 ± 6.1 16.3 ± 5.2

Band Squat depth (º) 95.2 ± 13.8 90.5 ± 21.7 102.3 ± 9.5 87.4 ± 12.5

(band: 0.79 ± 0.08, no band: 0.81 ± 0.07) (p=0.482). Similarly, KWI was not affected by the band during the eccentric phase of BW squats in the trained (band: 0.76 ± 0.08, no band: 0.73 ± 0.08) or the untrained group (band: 0.77 ± 0.70, no band: 0.75 ± 0.13) (p=0.670). However, during the concentric phase, there was a main effect of squat type (3RM or BW) with an overall lower KWI for 3RM squats (Figure 3) (p=0.046). Muscle Activity There was a significant main effect of the band condition, with use of the band increasing muscle activity in the majority of muscles during the eccentric (LVL, p=<0.001; RVL, p=<0.001; LBF, p=0.001; RBF, p=0.048; LGMe, p=0.001; RGMe, p=<0.001; LGMa, p=<0.001; RGMa p=<0.001; Table 2) as well as the concentric phase (LVL, p=<0.001; RVL, p=<0.001; LBF, p=0.044; RBF, p=0.025; LGMe,

Weight (kg) 129 ± 19.9 104 ± 8.9 83.2 ± 9.0 76.0 ± 8.9

Reps 3.0 3.0 15.5 ± 6.3 15.3 ± 10.4

Squat depth (º) 104.5 ± 18.2 87.9 ± 12.5 101.9 ± 12.8 91.7 ± 14.9

p=0.038; RGMe, p=0.049; RGMa, p=0.017; Table 3). For select muscles, a squat type by band interaction (LVL, p=0.019; LBF, p=0.035; LGMe, p= 0.020) was found for the eccentric phase of movement. Only LGMa (p=0.035) showed a training level X band interaction for the concentric phase of movement. Bonferoni post-hoc analysis showed differential results; while most muscles had increased muscle activity with band use BW squats, LVL demonstrated that band squats elicited lower EMG activity overall and this was more pronounced in the BW condition (p=0.026) regardless of phase (concentric vs eccentric). During BW squats, the LVL had significantly lower maximum muscle activity when performed with the band, regardless of trained (band: 115.03 ± 36.66% MVC, no band: 130.47 ± 44.54% MVC or untrained (band: 99.74 ± 53.61% MVC, no band: 118.09 ± 36.28% MVC) status.

Figure 3. Peak Knee Width Index (mean ± SD) for the trained group during the eccentric (A) and concentric (C) phases. Peak Knee Width Index (mean ± SD) for the untrained group during the eccentric (B) and concentric (D) phases. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 555

Table 2. Normalized maximum muscle activity (mean ± SD) for each muscle during the eccentric phase of all conditions (band, no band) and sessions (3RM, BW). Shaded bars represent trained group. *denotes a main effect of band (p<0.05) for that variable and ^denotes a signiﬁcant (p<0.05) interaction effect.

Table 3. Normalized maximum muscle activity (mean ± SD) for each muscle during the concentric phase of all conditions (band, no band) and sessions (3RM, BW). Shaded bars represent trained group. *denotes a main effect of band (p<0.05) for that variable and ^denotes a signiﬁcant (p<0.05) interaction effect.

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DISCUSSION This study investigated the effects of a resistance loop band, placed around the distal thigh, on medial knee collapse and muscle activity during the barbell back squat. More specifically, the band was evaluated in regards to training status (trained or untrained) and load (3RM or BW). Interestingly, there was a significant effect of load intensity (3RM or BW) on KWI, but no effect of band or training level conditions. Somewhat conversely, for the majority of muscles monitored, there was significantly greater muscle activity during the band conditions than no band conditions and this was not specific to training status or load. KWI, the primary measure of medial knee collapse for this investigation, showed no significant difference with respect to a band intervention. While many strength coaches indicate that the band helps promote a more neutral knee alignment and prevent medial knee collapse, the results of this study showed no significant effect of the band intervention, regardless of training status. This is in agreement with Gooyers et al.,12 who found similar conclusions during a bodyweight squat exercise. As suggested previously by Gooyers et al.,12 a longerterm intervention (with an untrained group) using a band may result in plastic changes to squat mechanics and performance, however a single set, regardless of load (3RM or BW), showed no change with the band in the current study. During the concentric phase of the squat, there was a main effect of squat type (3RM or BW) with an overall lower KWI for 3RM squats. This could be interpreted in two ways. One, the increased mechanical demand on the lower extremity due to greater squat load likely contributed to the lower KWI (more medial collapse) and, two, demands associated with the 3RM may have resulted in muscle fatigue, resulting in a lower KWI. Even though neither group showed improvements in KWI during the band loop conditions compared to no band, there is likely still room for improvement in KWI, possibly through the use of a longer training exposure using the band. Despite few changes in KWI during the 3RM or BW squat, the placement of a resistance band around the distal thighs did increase lower extremity muscle activity in both trained and untrained partici-

pants when compared to using no band. The band affected peak muscle activity for muscles in both phases of movement (eccentric: LVL, RVL, LBF, RBF, LGMe, RGMe, LGMa, RGMa; concentric phase: LVL, RVL, LBF, RBF, LGMe, RGMe, RGMa). The most consistent change across training level was for VL, which demonstrated consistently greater muscle activity with the band, across both the 3RM and BW conditions. For example, for the trained group, LVL had activity of 156.6, 168.7, 115.0 and 130.5 %MVC compared to 103.2, 142.2, 99.7, 118.1 %MVC for the untrained group during the 3RM band, 3RM no band, BW band and BW no band conditions, respectively. While Gooyers et al.12 did not measure muscle activity during their body weight squat exercise, the current findings support their hypothesis that the band may differentially change muscle activity patterns even though no reduction in medial collapse was quantified when using the band. Other studies have found no difference in hip abductor muscle activity when a resistance band is applied to a hip eccentric exercise.19 This is contradictory to the findings of this investigation, which demonstrate that the band increased activity across groups and conditions (effect of p<0.001 for eccentric: LVL, RVL, RGMe, LGMa, RGMa; concentric: LVL, RVL). This difference could be the results of a strategic difference when performing a heavy barbell squat as opposed to bodyweight hip centric exercises where smaller stabilizers can assume the roll of band resistance while the agonist muscles do not alter activity. Spracklin et al. 20, also demonstrated that the use of a looped resistance band increases hip muscle activity during a barbell back squat. The authors also concluded that squat performance (measured as number of repetitions completed) were not affected by the band. We also demonstrate no change in the number of repetitions completed, but also show few differences in lower extremity kinematics with respect to a band intervention. The band increased gluteal muscle activity (GMe and GMa) but only with consistency in the untrained participants. Trained participants demonstrated no difference in peak gluteal muscle activity (with the exception of LGMa in the 3RM squat) when using the band. It was hypothesized that only untrained participants would benefit from the resistance band preferentially, based on the premise that the trained

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participants have already achieved the muscle activation patterns required to promote neutral knee alignment and to resist medial collapse. The GMa and GMe are the major hip abductors acting to counter the band pull across the knees, but also the major initiator of hip extension, so it may be plausible that trained participants, when told to ‘keep the band tight’ are actually hindering their practiced and efficient gluteal activation patterns.19 The use of resistance band training aids is usually employed by coaches to benefit novice squatters, and these results confirm that trained groups would probably see little benefit. There are a few limitations that should be discussed. The trained group had a mean 3RM weight that was 125% greater than the untrained group. The authors did not explicitly aim for weight categories in this work, only focusing on training frequency and a minimum load of bodyweight. It is likely that there may not have been a large enough disparity between our trained and untrained groups. Because of the study design, participants in the untrained group had to squat their bodyweight. This means fairly active and strong males were recruited, and a group of lesser-trained participants might have benefited more from the band intervention. Both our groups presented with KWI ratios close to 1, having little room for improvement with the band intervention. Gooyers et al.,12 demonstrated that at maximum squat depth participants achieved a KWI of almost 1.0. This was for a body weight squat and demonstrates that the increased load during a barbell squat could negatively effect KWI. Additionally, it was deemed unacceptable to allow the untrained group to perform a 3RM with the band, prior to a regular unaided squat. Therefore, experimental days were not randomized and the trained group followed the same timeline for parity. The only cue given during the band intervention was ‘keep the band tight’. Perhaps with additional instruction, further differences would have been found in KWI and hip abductor activity with participants knowing when to activate abductors and extensors. Additionally, only one level of resistance band was used during this study. It is possible that using a gradient of resistances could have demonstrated a directional effect and it is possible that using heavier resistance would have elicited a greater response in peak muscle activation. Finally, it may be of interest to measure medial col-

lapse under load in a general population. Even being able to barbell back squat one’s own weight (the untrained group) requires a higher level of training than the greater population. CONCLUSIONS This study reports the first data on the neuromechanics of the lower limb during a barbell back squat with and without a band loop in both trained and untrained individuals. Squatting with a band increases lower limb muscle activation but does not change knee width index, and these observations were not training status or load dependent. Specifically, this work suggests that there is no evidence that a one-time exposure to a resistance band training aid changes the biomechanics of a barbell back squat as determined by knee width index, kinematics and muscle activity. Furthermore, this null result with respect to resistance band efficacy in reducing medial knee collapse was consistent across training experience. Despite little change in KWI for the untrained group, increased lower extremity muscle activity may suggest a more effective training paradigm for this population. Knee valgus, if left uncorrected, could lead to a variety of knee injuries, which has been shown to occur in a variety of sports that require squatting, explosive force and quick directional changes. Thus, it is critical to find ways to correct knee valgus and avoid these injuries. Future studies should incorporate different loop band resistances and incorporate both male and female participants. In addition, calculating knee collapse via changes in frontal plane kinematics and knee moments could provide further interpretation. This knowledge could, in turn, be used to develop training programs for athletes and rehabilitation patients who require a reduction in knee valgus loading in order to avoid potential injury or re-injury. REFERENCES 1. Bell DR, Vesci BJ, DiStefano LJ, Guskiewicz KM, Hirth CJ, Padua DA. Muscle Activity and Flexibility in Individuals With Medial Knee Displacement During the Overhead Squat. Athl Train Sport Health Care. 2012;4(3):117–125. 2. Claiborne TL, Armstrong CW, Gandhi V, Pincivero DM. Relationship between hip and knee strength

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and knee valgus during a single leg squat. J Appl Biomech. 2006;22(1):41–50. Holden S, Boreham C, Doherty C, Delahunt E. Two-dimensional knee valgus displacement as a predictor of patellofemoral pain in adolescent females. Scand J Med Sci Sport. (2015). doi: 10.1111/ sms.12633. Ford KR, Myer GD, Hewett TE. Valgus knee motion during landing in high school female and male basketball players. Med Sci Sports Exerc. 2003;35(10):1745–1750. Faul F, Erdfelder E, Lang A-G, Buchner A. G*Power 3: a ﬂexible statistical power analysis program for the social, behavioral, and biomedical Sciences. Behav Res Methods. 2007;39:175–91. Faul F, Erdfelder E, Buchner A, Lang A-G. Statistical power analyses using G*Power 3.1: tests for correlation and regression analyses. Behav Res Methods. 2009;41:1149–60. Gooyers CE, Beach TA, Frost DM, Callaghan JP. The inﬂuence of resistance bands on frontal plane knee mechanics during body-weight squat and vertical jump movements. Sport Biomech. 2012;11(3):438–439. Myer GD, Ford KR, Mclean SG, Hewett TE. The Effects of Plyometric Versus Dynamic Stabilization and Balance Training on Lower Extremety Biomechanics. Am J Sport Med. 2006;34(3):445–455. Geiser CF, O’Connor KM, Earl JE. Effects of isolated hip abductor fatigue on frontal plane knee mechanics. Med Sci Sports Exerc. 2010;42(3):535–545. Alenezi F, Herrington L, Jones P, Jones R. The reliability of biomechanical variables collected during single leg squat and landing tasks. J Electromyogr Kinesiol. 2014;24(5):718–721 Kritz M, Cronin J, Hume P. The Bodyweight Squat: A Movement Screen for the Squat Pattern. Strength Cond J. 2009;31(1):76–85. Hewett TE, Myer GD, Ford KR, Heidt RS, Colosimo AJ, McLean SG, van den Bogert AJ, Paterno MV, Succop P. Biomechanical Measures of

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Neuromuscular Control and Valgus Loading of the Knee Predict Anterior Cruciate Ligament Injury Risk in Female Athletes: A Prospective Study. Am J Sports Med. 2005;33(4):492–501. Hollman JH, Galardi CM, Lin IH, Voth BC, Whitmarsh CL. Frontal and transverse plane hip kinematics and gluteus maximus recruitment correlate with frontal plane knee kinematics during single-leg squat tests in women. Clin Biomech. 2014;29(4):468–474. More RC, Karras BT, Neiman R, Fritschy D, Woo SL, Daniel DM. Hamstrings--an anterior cruciate ligament protagonist. An in vitro study. Am J Sports Med. 1993;21(2):231–237. Lloyd DG, Buchanan TS. Strategies of muscular support of varus and valgus isometric loads at the human knee. J Biomech. 2011;34(10):1257–1267. Macrum E, Bell DR, Boling M, Lewek M, Padua D. Effect of limiting ankle-dorsiﬂexion range of motion on lower extremity kinematics and muscle-activation patterns during a squat. J Sport Rehabil. 2012;21(2):144–150. Hermens HJ, Freriks B, Disselhorst-Klug C, Rau G. Development of recommendations for SEMG sensors and sensor placement procedures. J Electromyography Kinesiolology. 2000;10:361–384. Noyes FR, Barber-westin SD, Fleckenstein C, Walsh C, West J. The Drop-Jump Screening Test Difference in Lower Limb Control By Gender and Effect of Neuromuscular Training in Female Athletes. Am J Sort Med. 2005;33(2):197-207. Frost DM, Beach T, Fenwick C, Callaghan J, Mcgill S. Is there a low-back cost to hip-centric exercise? Quantifying the lumbar spine joint compression and shear forces during movements used to overload the hips. J Sports Sci. 2012;30(9):859-870. Spracklin KF, Button DC, Halprin I. Looped band placed around thighs increases EMG of gluteal muscles without hindering performance during squatting. J Perf Health Res. 2017;1(1):60-71.

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IJSPT

ORIGINAL RESEARCH

A BIOMECHANICAL COMPARISON AMONG THREE KINDS OF REBOUND-TYPE JUMPS IN FEMALE COLLEGIATE ATHLETES Miki Nariai, MS1 Naruto Yoshida, PhD2 Atsushi Imai, PhD3 Kazumichi Ae, PhD4 Ryo Ogaki, PhD4 Hirokazu Suhara, MS1 Hitoshi Shiraki5

ABSTRACT Background: Single-legged drop jumps (SDJ), single-legged repetitive jumps (SRJ), and single-legged side hops (SSH) are often used as plyometric training and functional performance tests. Differences in the kinetics and kinematic characteristics of lower extremity joints during these jumps are unclear. Hypothesis/Purpose: The purpose of this study was to investigate the joint motion and mechanical work of the takeoff leg from foot contact to foot-off during SDJ, SRJ, and SSH in the sagittal and frontal planes in female athletes. It was hypothesized that the joint motion and mechanical work of the lower extremity joints during the SDJ and SRJ would be larger than the SSH in the sagittal plane, those during the SSH would be larger than the SDJ and SRJ in the frontal plane, and during SRJ would be larger than SDJ. Study Design: Cross-sectional study. Methods: Seventeen female collegiate athletes participated and performed the SDJ (0.15-m box height), and SRJ and SSH (by crossing two lines 0.3 m apart). Three-dimensional coordinate data and ground reaction forces were collected. Contact time, jump height, jump index (i.e., the jump height divided by the contact time) of the SDJ and SRJ, and the total times of the SSH were calculated. Range of motion (ROM) from touchdown to the lowest center of mass, and the positive and negative (mechanical) work from touchdown to foot-off were analyzed. Results: There were no significant differences in jump performance variables. Compared to the SSH, the SDJ and SRJ had significantly larger ankle and knee ROM and positive and negative work at the lower extremity joints, except for positive work at the hip joint, in the sagittal plane (p < 0.05). Compared to the SDJ and SRJ, the SSH had a significantly larger ankle ROM and positive work at the knee joint in the frontal plane (p < 0.05). Compared to the SDJ, the SRJ had a significantly larger ROM and negative work at each lower extremity joint in the frontal plane (p < 0.05). Conclusion: Although there were no significant differences in the jump performance variables, different characteristics of the takeoff leg ROM and mechanical work were found between three kinds of rebound-type jump tests. These findings may help clinicians choose jump methods to assess lower extremity function and to design plyometric training programs in sports and clinical fields. Level of Evidence: 3b Key words: mechanical work, plyometrics, three-dimensional motion analysis

Graduate School of Comprehensive Human Sciences, University of Tsukuba, Ibaraki, Japan. 2 Department of Acupuncture and Moxibustion, Faculty of Health Care, Teikyo Heisei University, Tokyo, Japan. 3 Faculty of Sport Sciences, Waseda University, Tokorozawa, Saitama, Japan. 4 Sports Research & Development Core, University of Tsukuba, Ibaraki, Japan. 5 Faculty of Health and Sport Sciences, University of Tsukuba, Ibaraki, Japan.

CORRESPONDING AUTHOR Miki Nariai, MS Graduate School of Comprehensive Human Sciences, University of Tsukuba, 1-1-1 Tennodai, Tsukuba, Ibaraki 305-8574, Japan E-mail: m.nariai1116@gmail.com

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INTRODUCTION Rebound-type jumps such as drop jumps (DJ) and repetitive jumps (RJ) are often used as plyometric training,1-2 injury predictors,3-5 and performance tests to assess the ability to jump.6-7 Each of these tasks makes use of the stretch shortening cycle (SSC), which is described as an eccentric contraction followed by an immediate concentric contraction of the muscle.8-9 For excellent sports performance, it is important to perform these jumps as high as possible, as well as quickly, with a short amount of ground contact time.10-11 Thus, the reactive strength index12 and RJ-index7,13 of the jump height divided by the contact time can be used assessment factors during DJ and RJ. As an assessment of lower extremity function related to knee injuries, previous authors have reported on the ability to perform single-legged DJ and RJ (SDJ and SRJ) in subjects with a history of knee injuries such as anterior cruciate ligament (ACL) injury.14-16 Moreover, Hewett et al. demonstrated that knee valgus motion and valgus moment during the double-legged DJ (DDJ) are predictors of ACL injury risk in female athletes.3 In previous studies about kinetic and kinematics in the sagittal plane during rebound-type jumps, it has been reported that the contribution of the ankles to perform DDJ with shorter contact time is larger than that to perform counter movement jumps,17-18 and Wang et al. showed that the peak power value of the ankle joint was larger than those of the knee and hip joint during takeoff in SDJ with 0.2-0.5 m box height.19 The results of these studies suggest that the ankles contribute to SDJ and SRJ performance. Yoshida et al. demonstrated that functional ankle instability (FAI) affects the contact time of the SDJ.20 Thus, the SDJ and SRJ test may be useful not only for assessing the capability in athletes with a history of the knee injury but also in those with a history of ankle injury. Regarding functional performance tests to assess ankle function, single-legged side hops (SSH) are often used. Previous studies have reported that the SSH time calculated by crossing two lines 0.3 m apart was associated with functional deficit in participants with FAI.21,22 Several researchers have investigated ankle joint motion, moments, and muscle activities during each phase of SSH. These authors demon-

strated larger inversion torque23 and muscle activity of the peroneus longus muscle23,24 during lateral hop contact than during medial hops. Although SSH are effective to assess the function of participants with FAI, lateral (horizontal) movements increase the load on muscles and ligaments controlling lateral displacement of each lower extremity joint. On the contrary, SDJ and SRJ may show a lower load on the muscles and ligaments because these jumps are performed in the vertical direction. Besides the direction of the jump, the number of jumps may also be related to the kinetics and kinematics during each jump. In a previous study, Kariyama et al. reported that frontal plane hip joint negative and positive work during SRJ was higher than during DRJ. 25 It is assumed that the biomechanical parameters of the lower extremity joints differ widely depending on number of jumps and jumping directions. However, no studies have compared the characteristics of each lower extremity joint in SSH with that of SDJ or SRJ. Therefore, the purpose of this study was to investigate the joint motion and mechanical work of the takeoff leg from foot contact to foot-off during SDJ, SRJ, and SSH in the sagittal and frontal planes in female athletes. The following was hypothesized: (1) the range of motion (ROM) and mechanical work at the lower extremity joints during the SDJ and SRJ would be larger than the SSH in the sagittal plane, whereas those during SSH would be larger than the SDJ and SRJ in the frontal plane; and (2) SRJ would have a larger ROM and mechanical work than SDJ, if the same drop height is used. METHODS Participants Seventeen female collegiate athletes (13 basketball players, 4 handball players) volunteered to participate in this cross-sectional study. Demographic data are presented in Table 1. None of the participants had any lower extremity injury or balance deficits caused by inner ear infection, cold, and taking medicine that would prevent them from completing all the trials. The Ethics Committee of the Faculty of Health and Sports Science at the University of Tsukuba approved this study, and all participants provided written informed consent before participating.

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Table 1. Demographic data of the participants

Participants (n) Age (y) Body height (m) Body mass (kg)

17 19.4 ± 1.4 1.67 ± 0.11 63.6 ± 12.6

Measurements and Procedures Participants performed a self-selected warm up for at least 10 minutes including dynamic movements such as jogging, stationary cycling, and jumping, among others. Then they performed the following jump performance tests: SDJ, SRJ, and SSH. All trials were performed with the participants’ dominant leg, which was defined as the takeoff leg when performing a running long jump. For the SDJ, participants performed a single-legged vertical drop jump (Figure 1a). The box height was set to 0.15 m, based on that used in a previous study.20 Participants were instructed to shorten the contact time as much as possible and jump as high as possible. Three successful trials were measured. For the SRJ,13 participants performed a single-legged repetitive rebound jump, which involved repeating the vertical jump five times (Figure 1b), upward off of the ground from a standing posture. Participants were given the same instructions for the SDJ, and then two successful trials were evaluated. In both tests, their hands were on their hips to counter the effect of arm swinging during the jumping performance7 and were not allowed to touch the floor with the opposite leg. For the SSH,21-22,24 participants performed a lateral hop over a 0.3 m distance, and then they returned to the starting position for 10 repetitions without stepping on the line with their test leg and touching the floor

with the opposite leg (Figure 1c); two successful trials were measured. Participants performed the tasks in the following order: SDJ, SRJ, and SSH. They had enough rest time between trials to avoid fatigue and ensure maximal performance on each trial. Data Collection and Data Processing Based on a study by Suzuki et al,26 three-dimensional coordinate data was collected using 47 retro-reflective markers attached to the body of each participant using a 10-camera motion capture system (250 Hz, VICON MX+; Vicon Motion Systems, Ltd., Oxford, UK). Data were smoothed using a Butterworth digital filter with no phase lags set at optimal cutoff frequencies (7.5– 15 Hz).27 Next, the ground reaction force data were recorded using a force platform (9287B, 1000 Hz; Kistler Instruments AG, Winterthur, Switzerland). Threedimensional coordinate data were spline-interpolated to 1000 Hz to calculate kinetic parameters combined with the ground reaction force data. Data Analysis Jump performance variables To assess each participant’s jump performance ability, the jump height (m), contact time (s), and jump index were measured for the SDJ and SRJ. The jump heights during SDJ and SRJ were calculated as follows: JH = 1/8 ⫻ g ⫻ FT2 where JH is the jump height, g is gravitational acceleration (m/s2), and FT is the flight time (s).28 The jump index, which was the ratio of the jump height

Figure 1. Three kinds of jump performance tests. a: single-legged rebound drop jump, b: single-legged repetitive jump, c: single-legged side hop. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 562

divided by the contact time, was based on the reactive strength index12 and RJ-index.13 The highest indeces were selected for further analysis. In the SDJ, the trial resulting in the highest index of the three trials was adopted and biomechanical analysis was done. For the SRJ, the jump index was calculated for all jumps of the two trials and the jump with the highest jump index was further analyzed. For the SSH, the total time (s) of ten repetitions was measured and the fastest time of the two trials was adopted. In the selected trial, contact time (s) of fifth lateral contact was calculated and biomedical analysis of this jump was performed. ROM and mechanical work The contact phase from the touchdown to foot off was used as the analysis period for the ROM and mechanical work. The eccentric phase was defined from the touchdown to the lowest point of the center of gravity in the contact phase. The ROM and mechanical work of the takeoff leg joints were calculated in the sagittal and frontal planes: sagittal plane, dorsiflexion/plantar flexion axis at the ankle joint, flexion/extension axis at the knee joint, flexion/extension at the hip joint axis, frontal plane and eversion/inversion axis at the ankle joint, varus/valgus axis at the knee joint, and adduction/abduction axis at the hip joint. The ROM was determined from the maximum to minimum joint angle at each joint, which was calculated during the eccentric phase. Mechanical work during the contact phase was calculated by integrating the joint torque power.29 The mechanical work was normalized by each participant’s body weight. In this study, positive work represented energy generation, whereas negative work represented energy absorption. The mass, center of gravity, and moments of inertia were calculated for each segment using body segment coefficients provided by Ae.30 Data analysis was conducted using

MATLAB (The MathWorks, Inc., Natick, MA, USA). The coordinate system was defined according to a previous study by Kariyama et al.31 Statistical Methods All data are presented as a mean ± standard deviation. One-way repeated-measure analysis of variance with post-hoc Bonferroni corrections was used to assess mean differences in the contact time, ROM, and mechanical work among the SDJ, SRJ, and SSH. A paired-sample t-test was used to examine mean differences in the jump height and jump index between the SDJ and SRJ. Statistical analyses were performed using the IBM SPSS statistics, version 21 package (IBM Corp., Armonk, NY, USA), and the significance level was set at p < 0.05. RESULTS Performance Variable Data for the contact time, jump height, and jump index for each jump performance test are presented in Table 2. There were no significant differences in any variables among the jump performance tests (p > 0.05). ROM Table 3 shows the ROM of the ankle, knee, and hip joints in the eccentric phase. In the sagittal plane, the ankle ROMs were the largest and smallest during the SRJ and SSH, respectively (p < 0.05). The knee ROMs for the SDJ and SRJ were significantly larger than that demonstrated during the SSH (p < 0.05). In the frontal plane, the ankle ROM were the largest and smallest during the SSH and SDJ, respectively (p < 0.05). The knee ROMs for the SRJ and SSH were significantly greater than that of the SDJ. The hip ROMs for the SRJ was significantly larger than that of the SDJ (p < 0.05), and there were no significant differences between SSH and SDJ and SRJ.

Table 2. Results of the jump performance tests

Contact time (s) Jump height (m) Jump index (m/s) Total time (s)

SDJ 0.246 ± 0.025 0.127 ± 0.030 0.525 ± 0.148

SRJ 0.247 ± 0.026 0.133 ± 0.022 0.548 ± 0.130

SSH 0.231 ± 0.031

7.713 ± 0.815

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Table 3. Range of motion of the ankle, knee, and hip joints

Sagittal plane

Frontal plane

SDJ

SRJ

SSH

Ankle (°)

34.83 ± 5.41*†

40.95 ± 3.79*‡

21.90 ± 6.49†‡

Knee (°)

25.91 ± 5.22†

22.95 ± 4.39‡

12.78 ± 5.60†‡

Hip (°)

6.27 ± 2.52

7.64 ± 2.33

7.74 ± 2.61

Ankle (°)

5.67 ± 3.31*†

8.88 ± 2.51*‡

14.20 ± 4.63†‡

Knee (°)

4.91 ± 1.06*†

6.60 ± 1.49*

7.46 ± 2.05†

Hip (°)

4.21 ± 2.31*

5.95 ± 2.89*

4.60 ± 2.09

SDJ, single-legged rebound drop jumps; SRJ, single-legged repetitive jumps; SSH, single-legged side hops *

Significant difference between the SDJ and SRJ (p < 0.05).

†

Significant difference between the SDJ and SSH (p < 0.05).

‡

Significant difference between the SRJ and SSH (p < 0.05).

Table 4. Mechanical work in the sagittal plane A nk le dorsiflexion/plantarflexion SDJ SRJ SSH

Knee flexion/extension SRJ

Hip flexion/extension SRJ

SDJ SSH SDJ SSH Positive work 1.31 ± 0.14† 0.74 ± 0.23† 0.70 ± 0.21‡ 0.35 ± 0.21†‡ 0.09 ± 0.06† 0.08 ± 0.06‡ 0.14 ± 0.07†‡ 1.37 ± 0.18‡ 0.59 ± 0.12†‡ (J/kg) Negative work -0.73 ± 0.25† -0.73 ± 0.22‡ -0.29 ± 0.14†‡ -0.20 ± 0.10† -0.16 ± 0.08‡ -0.09 ±0.05†‡ -1.18 ± 0.21*† -1.39 ± 0.28*‡ -0.67 ± 0.23†‡ (J/kg) SDJ, single-legged rebound drop jumps; SRJ, single-legged repetitive jumps; SSH, single-legged side hops * Significant difference between the SDJ and SRJ (p < 0.05). † Significant difference between the SDJ and SSH (p < 0.05). ‡ Significant difference between the SRJ and SSH (p < 0.05).

Mechanical Work Table 4 shows the mechanical work in the sagittal plane. The positive work at the ankle and knee for the SDJ and SRJ was significantly greater than that during the SSH (p < 0.05). The positive work at the hip joint for the SSH was greater than that for the SDJ and SRJ. The negative work at the ankle joint were the greatest and least during the SRJ and SSH, respectively (p < 0.05). In addition, the negative work at the knee and hip joints of the SDJ and SRJ was significantly greater than that during the SSH (p < 0.05). Table 5 shows the mechanical work in the frontal plane. The positive work at the knee were the greatest and least during the SSH and SDJ,

respectively (p < 0.05). The negative work at each lower extremity joint for the SRJ and SSH was significantly greater than that during the SDJ (p < 0.05). DISCUSSION The main results showed that the SRJ and SDJ resulted in larger ROM and mechanical work at the ankle and knee in the sagittal plane than the SSH, whereas the SSH resulted in larger ankle eversion/ inversion ROM and positive work at the knee varus/ valgus axis and the hip flexion/extension axis than the SDJ and SRJ. Moreover, the SRJ had larger ROM and negative work at each lower extremity joint in the frontal plane than the SDJ.

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Table 5. Mechanical work in the frontal plane

Positive work (J/kg) Negative work (J/kg)

SDJ

Ankle inversion/eversion SRJ

SDJ

Knee varus/valgus SRJ

SSH

0.01 ± 0.01

0.02 ± 0.01

-0.02 ± 0.01*†

-0.03 ± 0.01*

Hip adduction/abduction SRJ

SSH

SDJ

0.03 ± 0.03

0.08 ± 0.04*†

0.10 ± 0.05*‡

0.15 ± 0.08†‡

0.25 ± 0.11

0.26 ± 0.12

0.21 ± 0.07

-0.04 ± 0.03†

-0.08 ± 0.04*†

-0.11 ± 0.05*

-0.15 ± 0.06†

-0.14 ± 0.09*†

-0.21 ±0.13*

-0.18 ± 0.10†

SSH

SDJ, single-legged rebound drop jumps; SRJ, single-legged repetitive jumps; SSH, single-legged side hops * Significant difference between the SDJ and SRJ (p < 0.05). † Significant difference between the SDJ and SSH (p < 0.05). ‡ Significant difference between the SRJ and SSH (p < 0.05).

Jump Performance Variables There were no significant differences in any jump performance variables among the jump performance tests. The SSC exercises is categorized as either fast or slow depending on the ground contact times.32,33 These jumps may be categorized as fast SSC movement, because all three kinds of jumps had similar contact times <0.250 s. Moreover, participants were able to perform at a similar ability during these jumps since the jump height and jump index were not significantly different between the SDJ and SRJ. ROM and Mechanical Work First, the SDJ and SRJ was compared with the SSH to determine differences between rebound-type jumps in the vertical and lateral directions. In the sagittal plane, it was found that the SDJ and SRJ had a larger ROM and mechanical work around the ankle dorsiflexion/plantar flexion axis and knee flexion/extension axis (Table 4). Therefore, these results partially support the hypothesis that the ROM and mechanical work at each lower extremity joint during the SDJ and SRJ would be larger than the SSH in the sagittal plane larger than the SSH during the SDJ and SRJ in the sagittal plane. Moreover, the SSH had the largest amount of positive work around the hip flexion/extension axis among the three jumps (Table 4). This observation was consistent with those from a previous study that investigated contributions to lateral propulsion.34 Inaba et al. reported that positive work around the hip flexion/extension axis contributed to lateral propulsion in the medial direction during side-step takeoff.34 The current study found that the SDJ and SRJ, which are primarily vertical movements, had larger joint motion during the

eccentric phase and mechanical work during takeoff at the ankle and knee than the SSH. In addition, more hip flexion/extension energy generation may be required to jump in the medial direction during the SSH than in the vertical direction during the SDJ and SRJ. Compared to the SSH and SDJ, the SSH had a larger ankle eversion/inversion and knee varus/valgus ROM (Table 3), and the SSH had a larger negative work at each lower extremity joint compared to the SDJ in the frontal plane (Table 5). These results partly support the hypothesis that the ROM and mechanical work at each lower extremity joint would be larger for the SSH than those for the SDJ and SRJ in the frontal plane. However, when comparing the SSH to the SRJ, almost all variables were not significantly different. It was expected that the foot contact position would be easier to displace slightly laterally during the SRJ than during the SDJ. Since participants may have been required to correct the trunk and lower extremity joint positions in SRJ and SSH, it was considered that they have similar joint motion and mechanical work in the frontal plane. Second, the differences between the SDJ and SRJ, which were performed in the same direction, were assessed. The current results showed that the SRJ had a larger ankle dorsiflexion/plantar flexion ROM (Table 3) and negative work around the ankle dorsiflexion/plantar flexion axis (Table 4). Additionally, negative work around each joint axis in the frontal plane (Table 5) was larger than that during the SDJ. Previous authors have reported that the ground reaction force, as well as the kinetics and kinematics of the lower extremity joint, were affected by the land-

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ing height.35-36 However, the jump height of the SRJ (0.133 ± 0.022 m) was similar to the box height of the SDJ (0.15 m) in the current study. Hence, it seems that there is no need to consider the effect of the drop height as a confounding variable. However, the number of jumps is different between the SDJ and SRJ. Because the SRJ requires repeated absorption impact forces and continual postural control, these characteristics of the SRJ may cause an increase in ROM and negative work. Practical Application Clinicians often use SDJ, SRJ, and SSH to assess or improve lower extremity function. The results of the current study indicate that SDJ, SRJ, and SSH have similar contact times, but they demonstrated differences in ROM and mechanical work of the takeoff leg. These results may provide helpful information so clinicians can choose the best jump method to assess the function of the lower extremity and plan plyometric training programs according to the subjects’ history of injury and level of functional deficit. In the current study, the SRJ had more negative work at the ankle joint in the sagittal and frontal planes than the SDJ; thus, it is considered that the SRJ causes more load and stress on the ankle joint than the SDJ with a 0.15-m box height. SDJ may be useful as a functional performance test before the beginning of the exercise, as well as SRJ, which requires a jump height greater than 0.15 m, and SSH. Moreover, previous authors have demonstrated the effect of knee injury on the ability to perform SRJ14-15 and the time factor during performance of the SSH.37 Accordingly, clinicians should remember that jump performance variables measured by these jump tasks may be affected by functional deficits of the knee or hip joints, in addition to those deficits present in the ankle. Limitations The kinetics and kinematics during rebound-type jumps we measured only in healthy female athletes. Thus, these characteristics in male athletes and athletes with a history of injury are not yet clear. Future studies should focus on athletes, including men and those with a history of lower extremity injury, as well as during other phases during the SSH in order

to clarify which types of jump performance tests and training could be used to assess and improve athletic performance. CONCLUSION Compared to the SDJ and SRJ, SSH had larger ankle eversion/inversion ROM and positive work at the knee varus/valgus axis and the hip flexion/extension axis. Moreover, the SRJ had larger ROM and negative work at each lower extremity joint in the frontal plane than the SDJ, although the two jumps are performed in the same direction. REFERENCES 1. Huang PY, Chen WL, Lin CF, et al. Lower extremity biomechanics in athletes with ankle instability after a 6-week integrated training program. J Athl Train. 2014;49:163-172. 2. Myer GD, Ford KR, Brent JL, et al. The effects of plyometric vs. dynamic stabilization and balance training on power, balance, and landing force in female athletes. J Strength Cond Res. 2006;20:345-353. 3. Hewett TE, Myer GD, Ford KR et al., Biomechanical measures of neuromuscular control and valgus loading of the knee predict anterior cruciate ligament injury risk in female athletes: a prospective study. Am J Sports Med. 2005;33:492-501. 4. Hewett TE, Myer GD, Kiefer AW, et al. Longitudinal increases in knee abduction Moments in Females during adolescent growth. Med Sci Sports Exerc. 2015;47:2579-2585. 5. Myer GD, Ford KR, Palumbo JP, et al. Neuromuscular training improves performance and lower-extremity biomechanics in female athletes. J Strength Cond Res. 2005;19:51-60. 6. Young W. Laboratory strength assessment of athletes. New Student Athlete. 1995;10:88-96. 7. Tauchi K, Endo T, Ogata M, et al. The characteristics of jump ability in elite adolescent athletes and healthy males. The development of countermovement and rebound jump ability. Int J Sport Health Sci. 2008;6:78-84. 8. Norman RW, Komi PV. Electromechanical delay in skeletal muscle under normal movement conditions. Acta Physiol Scand. 1979;106:241-248. 9. Komi PV. Stretch-shortening cycle: a powerful model to study normal and fatigued muscle. J Biomech. 2000;33:1197-1206. 10. Barr MJ, Nolte VW. Which measure of drop jump performance best predicts sprinting speed? J Strength Cond Res. 2011;25:1976-1982.

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11. Sattler T, Sekulic´ D, Spasic´ M, et al. Analysis of the association between motor and anthropometric variables with change of direction speed and reactive agility performance. J Hum Kinet. 2015;29:137-145. 12. Flanagan EP, Ebben WP, Jensen RL. Reliability of the reactive strength index and time to stabilization during depth jumps. J Strength Cond Res. 2008;22:1677-1682. 13. Miura K, Yamamoto M, Tamaki H, et al. Determinants of the abilities to jump higher and shorten the contact time in a running 1-legged vertical jump in basketball. J Strength Cond Res. 2010;24:201-206. 14. Flanagan EP, Galvin L, Harrison AJ. Force production and reactive strength capabilities after anterior cruciate ligament reconstruction. J Athl Train. 2008;43:249-257. 15. Schiltz M, Lehance C, Maquet D, et al. Explosive strength imbalances in professional basketball players. J Athl Train. 2009;44:39-47. 16. Willson JD, Davis IS. Utility of the frontal plane projection angle in females with patellofemoral pain. J Orthop Sports Phys Ther. 2008;38:606-615. 17. Bobbert MF, Huijing PA, and van Ingen Schenau GJ. Drop jumping. I. The influence of jumping technique on the biomechanics of jumping. Med Sci Sports Exerc. 1987;19:332-338. 18. Bobbert MF, Mackay M, Schinkelshoek D, et al. Biomechanical analysis of drop and countermovement jumps. Eur J Appl Physiol Occup Physiol. 1986;54:566-573. 19. Wang LI, Peng HT. Biomechanical comparisons of single- and double-legged drop jumps with changes in drop height. Int J Sports Med. 2014;35:522-527. 20. Yoshida N, Miyakawa S, Miyamoto T, et al. Effect of functional ankle instability on rebound drop jump. J Appl Phys Fitness Sports Med. 2012;1:679-684. 21. Caffrey E, Docherty CL, Schrader J et al. The ability of 4 single-limb hopping tests to detect functional performance deficits in individuals with functional ankle instability. J Orthop Sports Phys Ther. 2009;39:799-806. 22. Docherty CL, Arnold BL, Gansneder BM, et al. Functional-performance deficits in volunteers with functional ankle instability. J Athl Train. 2005;40:30-34. 23. Monteleone BJ, Ronsky JL, Meeuwisse WH, et al. Lateral hop movement assesses ankle dynamics and muscle activity. J Appl Biomech. 2012;28:215-221. 24. Yoshida M, Taniguchi K, Katayose M. Analysis of muscle activity and ankle joint movement during the side-hop test. J Strength Cond Res. 2011;25:22552264.

25. Kariyama Y, Fujii H, Mori K, et al. The comparison to the characteristics of three-dimensional joint kinetics between single leg and double legs rebound jump. Proceedings of the 30th Annual Conference of Biomechanics in Sports. 2012;25-28. 26. Suzuki Y, Ae M, Takenaka S, et al. Comparison of support leg kinetics between side-step and cross-step cutting techniques. Sports Biomech. 2014;13:144-153. 27. Wells RP, Winter DA. Assessment of signal and noise in the kinematics of normal, pathological and sporting gaits. In: Human Locomotion 1: (Proceedings of the First Biannual Conference of the Canadian Society of Biomechanics). 1980:92-93. 28. Bosco C, Luhtanen P, Komi PV. A simple method for measurement. of mechanical power in jumping. Eur J Appl Physiol Occup Physiol. 1983;50:273-282. 29. Winter DA. Moments of force and mechanical power in jogging. J Biomech. 1983;16:91-97. 30. Ae M. Body segment inertia parameters for Japanese children and athletes (in Japanese). Japanese J Sports Sci. 1996;15:155-162. 31. Kariyama Y, Hobara H, Zushi K. Differences in take-off leg kinetics between horizontal and vertical single-leg rebound jumps. Sports Biomech. 2016; 4:1-14. 32. Schmidtbleicher D. Training for power events. In: The Encyclopedia of Sports Medicine. Vol 3: Strength and Power in Sport. Oxford, UK: Blackwell, P.V Komi, ed.,1992;169-179. 33. Flanagan EP, Comyns TM. The use of contact time and the reactive strength index to optimize fast stretch-shortening cycle training. J Strength Cond Res. 2008;30:32-38. 34. Inaba Y, Yoshioka S, Iida Y, et al. A biomechanical study of side steps at different distances. J Appl Biomech. 2013;29:336-345. 35. Bobbert MF, Huijing PA, van Ingen Schenau GJ. Drop jumping. II. The influence of dropping height on the biomechanics of drop jumping. Med Sci Sports Exerc. 1987;19:339-46. 36. Peng HT. Changes in biomechanical properties during drop jumps of incremental height. J Strength Cond Res. 2011;25:2510-2518. 37. Itoh H, Kurosaka M, Yoshiya S, et al. Evaluation of functional deficits determined by four different hop tests in patients with anterior cruciate ligament deficiency. Knee Surg Sports Traumatol Arthrosc. 1998;6:241-245. 38. Kurabayashi J, Mochimaru M, Kouchi M. Validation of the estimation methods for the hip joint center (in Japanese). J Soc Biomech. 2003;27:29-36.

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APPENDIX 1 Deﬁnition of the coordinate system ﬁxed at the hip, knee, and ankle joints for calculating the joint angles and torque of each joint. Figure 2 shows the definition of the coordinate system at the hip, knee, and ankle joints for calculating the joint angles and torques. The centers of the knee and ankle joints and foot were set at the midpoints between the medial and lateral markers. The hip joint center was estimated using the method recommended by the clinical gait analysis forum of Japan.38 The following describes how the joint coordinate system of the right hip joint and the pelvic coordinate system were determined to calculate the hip flexion/extension angle: For the pelvic coordinate system, xpel was defined as a vector from the center of the left hip joint to the center of the right hip joint; zpel was defined as the cross product of xpel and the vector from the midpoint of the left and right posterior superior iliac spine to the midpoint of the left and right anterior superior iliac spine; and ypel was defined as the cross product of xpel and zpel. For the right hip flexion/extension coordinate system, zhFE was defined by the cross product of xhFE (xpel) and a vector from the center of the right knee joint to the center of the right hip joint; yhFE was defined by the cross product of zhFE and xhFE (Figure 2a). For the right hip adduction/ abduction coordinate system, xhAA was defined by the cross product of yhAA (yhFE) and zhAA (from the center of the right knee to the center of the right hip). The hip flexion/extension angle was defined as the angle between ypel and yhFE, and the hip adduction/abduction angle was defined as the angle between zhFE and zhAA.

Figure 2. Deﬁnition of the coordinate system for the lower extremity. 1: center of the left hip joint 2: center of the right hip joint, 3: medial aspect of right knee 4: lateral aspect of right knee, 5: center of the right knee joint, 6: right malleolus medialis, 7: right malleolus lateralis, 8: center of the right ankle joint, 9: 1st metatarsal bone of right foot, 10: 5th metatarsal bone of right foot, 11: center of right foot, 12: right heel. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 568

IJSPT

ORIGINAL RESEARCH

CADAVERIC EVALUATION OF THE LATERAL-ANTERIOR DRAWER TEST FOR EXAMINING POSTERIOR CRUCIATE LIGAMENT INTEGRITY Gesine H. Seeber, PT, MS1 Marc P. Wilhelm, PT, DPT2 Gunther Windisch, MD3 Hans-Joachim Appell Coriolano, MD4 Omer C. Matthijs, PT, ScD2 Philip S. Sizer, PT, PhD2

ABSTRACT Background: Common clinical tests often fail to identify posterior cruciate ligament (PCL) ruptures, leading to undetected tears and potential degenerative changes in the knee. The lateral-anterior drawer (LAD) test has been proposed but not yet evaluated regarding its effectiveness for diagnosing PCL-ruptures. Hypothesis: The LAD will show greater tibial translation values in lateral-anterior direction in a PCL-Cut condition compared to a PCL-Intact condition, thus serving as a useful test for clinical diagnosis of PCL integrity. Study Design: Descriptive laboratory study. Methods: Threaded markers were inserted into the distal femur and proximal tibia in eighteen cadaveric knees. Each femur was stabilized and the tibia translated in lateral-anterior direction for the LAD test versus in a straight posterior direction for the posterior sag sign (PSS). Each test was repeated three times with the PCL both intact and then cut, in that order. During each trial, digital images were captured at start and finish positions for the evaluation of tibial marker displacement. Tibial marker translation during each trial was digitally analyzed using photography. The PSS values served as a reference standard. Results: The LAD tibial translation was significantly greater (U=-3.680; p<0.002) during the PCL-Cut (10.6±5.6mm) versus PCL-Intact (7.7±5.1mm) conditions. The PSS tibial translation was significantly greater (U=-3.724; p<0.002) during the PCL-Cut (11.0±5.3mm) versus PCL-Intact (6.4±3.5mm) conditions. There was no significant difference (t=2.029; p=0.07) in mean tibial translation in respective directions after PCL dissection during the LAD test (2.9±2.1mm) versus the PSS (4.6±2.8mm). Conclusion: The LAD test detected changes in cadaveric tibial translation corresponding with changes in PCL integrity to a degree at least as effective for assessing PCL integrity as the PSS. Further clinical study will be required to assess the utility of the LAD as a physical examination tool for diagnosing PCL injuries. Level of Evidence: 2 (laboratory study) Key words: Lateral-anterior drawer test, posterior cruciate ligament, tibial translation

University Hospital for Orthopaedics and Trauma Surgery Pius-Hospital, Medical Campus Carl von Ossietzky University Oldenburg, Germany 2 Center for Rehabilitation Research, School of Health Professions, TTUHSC, Lubbock, TX, USA 3 Praxis f. Manuelle Medizin, Graz, Austria 4 Department of Physiology and Anatomy, German Sports University Cologne, Germany Financial Disclosure and Conflict of Interest: We affirm that we have no financial affiliation (including research funding) or involvement with any commercial organization that has a direct financial interest in any matter included in this manuscript.

CORRESPONDING AUTHOR Phillip S. Sizer Jr, PT, PhD Center for Rehabilitation Research Department of Rehabilitation Sciences School of Health Professions Texas Tech University Health Sciences Center 3601 4th Street,6280, Lubbock, Texas, USA E-mail: Phil.sizer@ttuhsc.edu 806-743-3902

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INTRODUCTION Approximately two million sports injuries are registered in Germany per year, and 30% of these cases involve the knee joint,1 making it the second most involved body region for sports injury at a national level.2 A posterior cruciate ligament (PCL) rupture is considered to be one of the more severe sportsrelated knee injuries.3 The epidemiology of PCL injuries has not been sufficiently clarified to date.4,5 Depending on the patient population, the reported incidence of PCL injuries ranges between 1% and 44% of all acute knee injuries.6-10 However, the estimated number of undetected PCL ruptures is high,11,12 thus, epidemiological data regarding actual incidence might even be greater. Moreover, detection of a PCL-rupture can be delayed or missed, often due to mild acute symptoms, poorly understood pathology, and insufficient diagnostic testing.3 A PCL-injury can be sustained during many sports including but not limited to snow skiing, martial arts, soccer, rugby, and American football.1,3,13,14 Overlooked or misdiagnosed PCL-ruptures can lead to knee joint biomechanical changes,15,16 ultimately lending to the development of irreversible secondary failures, and serious functional impairments.11,12,14,17 Undetected knee instability related to PCL-rupture can produce early cartilage degeneration in response to increased compression and shearing forces in the medial and patellofemoral knee joint compartments.12,16,17 In addition, PCL-insufficiency leads to collagen maladaptation and functional deficits similar to those seen in ACL insufficiency.3,8,18,19 Concomitant posterolateral corner injury can lead to additional knee instability and degenerative changes.20 Thus, PCL-insufficiency detection is critical, necessitating precision clinical testing. Three clinical tests are most commonly used for examining PCL integrity, including the posterior sag sign (PSS), the posterior drawer test and the quadriceps active test.21 However, these often do not clearly identify PCL-ruptures.3,22 Kopkow et al.23 used a systematic review to report the clinical accuracy (sensitivity, specificity, positive and negative likelihood ratios) of these tests and found that the interpretation may be invalid due to poor methodology in most studies that were included.23 The posterior drawer test examines the supine patient with the knee flexed

to 80-90 degrees and the foot flat on the table. The tibia is passively translated posteriorly, parallel to the tibial plateau. However, often the patient presents with a posteriorly subluxed tibia (at rest, prior to performing the test) due to gravity, lending to minimal or unavailable posterior passive tibial movement, producing a false negative test.3,12 Similarly, joint swelling or obesity could hinder an examinerâ&#x20AC;&#x2122;s ability to palpate anterior tibial step-off, possibly decreasing posterior drawer test clinical accuracy.12 Similar interpretation challenges arise with the interpretation of the PSS and/or the quadriceps active test, where test outcomes depend on consistent quadriceps muscle relaxation. Both tests are likely to be misinterpreted if the patient exhibits apprehension or increased post-traumatic muscle tone. Finally, the use of medical imaging for diagnosing PCL-ruptures should be viewed with caution.24 Because the PCL can heal in an elongated position21,25-27 it may appear intact while in fact remaining lax. The LAD test, which is an adaption of the lateral shear test of Cyriax, has been introduced as a manually applied testing alternative that attempts to resolve the previously discussed challenges encountered while testing for PCL-rupture.28 In contrast to other previously established clinical PCL tests, utilizing the LAD test does not require precise palpation of any bony landmark, any visual rating of tibial displacement, or considerable muscle relaxation, due to the testing direction that is out of any muscleâ&#x20AC;&#x2122;s functional plane. Additionally, utilizing the LAD test might be psychologically advantageous for the patient since the manual load is not applied in the direction of instability. Any comparison of the LAD test to other well-accepted tests for identifying PCL-rupture has yet to be conducted. Therefore, the purpose of this study was to establish whether the LAD test sufficiently detects changes in in-vitro tibial translation when the PCL is disrupted. This was achieved by examining intra-specimen alterations of lateral-anterior tibial translation in specially embalmed cadaveric knees during both intact- and cut-PCL conditions. Moreover, the authors further assessed the test utility by comparing the changes in LAD translation data produced during both PCL conditions versus changes in PSS tibial translation data produced during the same PCL conditions.

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METHODS Subjects The cadaveric specimen sample consisted of 18 previously embalmed right-sided cadaveric lower extremities (36-94 years old; mean 79 years) sectioned from pelvis to foot. The cadavers, embalmed according to the method of Thiel,29 fulfilled the following criteria: • A complete right leg with pelvis • Undamaged muscle tissue • No visible (surgical) scars in the knee region • Intact cruciate ligaments* • Undamaged posterior capsule* (*Verified through arthroscopy performed between PCL conditions) The Thiel cadaveric preservation method is known to sustain a high degree of suppleness, allowing near-natural passive movement of all body parts.29 Utilizing these naturally preserved cadavers enabled the investigator to perform both the LAD test and PSS under conditions similar to the movement conditions found in a living person.29 Pre-measurement preparation of the anatomical specimen Threaded markers (commercially available 2cm Phillips-head screws) were inserted into the distal femur and proximal tibia in order to compare the extent of tibial translation produced with the PCLIntact versus PCL-Cut during both the LAD and PSS testing procedures, respectively. Markers were placed in a standardized manner, positioning the knee in 90° flexion on a wooden wedge (37 x 50 x 32 cm3) that was fixed to the treatment table and the tibia positioned in neutral rotation (at the midpoint between maximum available tibial external and internal rotation). For the LAD test assessment, markers were inserted 3 and 5 cm distal to the knee joint line at the anterior-medial proximal tibia and 3 cm proximal to the knee joint line at the anteriormedial distal femur. For the PSS assessment, markers were inserted in the lateral proximal tibia 3 cm distal to the knee joint line, as well as in the lateral distal femur 3 cm proximal to the same joint line.

Instrumentation To objectively record tibial translation changes produced during the PCL-Intact versus PCL-Cut conditions during both the LAD and PSS tests, digital images were captured using a commercial-quality digital reflex camera (Nikon D 70s, Nikon Corp., Japan) mounted on a commercial-grade photo tripod (Manfrotto 322RC2, Manfrotto Distribution, Cassola, Italien). All joint positions during each experimental stage were verified with an inclinometer (Baseline Bubble Inclinometer; 3B Scientific, Fabrication Enterprises INC. White Plains, NY, USA). The fluidfilled inclinometer´s scale was set at zero in full knee extension prior to every usage. A hand-held dynamometer (MicroFet2TM; Hoggan Health Industries, West Jordan, UT, USA) was utilized during all LAD trials in order to insure consistent force reproduction during test procedures. The device was manually calibrated prior to every usage. Preparatory Procedures To ensure measurement validity the camera was aligned perpendicular to each testing plane, the focal distance was established and maintained through all measurements, and an object of known dimension was placed in a consistent position in the visual field (to be later used during the digitization process for measurement calibration purposes).30 To insure compliance with these criteria, a ruler of standard length (10-20 cm) was mounted on a wooden wedge in a consistent position with respect to the cadaveric specimen and camera, while the camera was positioned as follows prior to data capture: 1. For the LAD test-the camera was positioned 1 m away from the specimen along a line that was 45 degrees anterior-medial to the tibia, allowing for data capture in an oblique plane. 2. For the PSS-the camera was positioned 1 m away from the specimen, directly lateral to the tibia at the height of the joint line, allowing for data capture in the parasagittal plane. Data capture utilized a within-camera focal length of 55 mm in natural room illumination. The camera was manually focused during each trial prior to data capture. Digital images of marker start and finish positions were used to establish tibial marker translation for each condition during each test trial.

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The cadaveric leg was repeatedly moved in flexion and extension at the knee and hip joint for five minutes to reduce tissue stiffness and minimize any residual muscle resistance. The specimen was then placed on the wooden wedge as previously described. The bony pelvis was fixated with a belt to ensure stability. The screw markers were applied to the tibia and femur as previously described. Data Collection Procedures The leg was adjusted to the position for PSS (90° hip flexion, 90° knee flexion, calf free-floating, and heel stabilized on a wooden block), consistent with previous reports. When used in-vivo, this position aims to maximize the influence of gravity on posterior displacement.31 For the PSS start position the tibia was manually translated to the maximum anterior position (Figure 1). For the PSS finish position the tibia was allowed to fully sag posteriorly in response to gravitational pull (Figure 2). The PSS trial was repeated three times. Digital images of start and finish positions were captured for each trial. Following the third PSS trial the leg was repositioned for the LAD test. The LAD test examines the knee while positioned in 90° flexion and the hip positioned at 45° flexion. While being aligned rather oblique to the tibial plateau in the extended knee, the PCL describes a circular arc of more than 60° during knee flexion,32 thus lending the PCL alignment to become gradually more vertical oriented with respect to the tibial plateau, attaining a near perpendicular posi-

Figure 1. Start position of the PSS; T = Tibia; F = Femur; L = Line coursing parallel to tibial diaphysis through lateral tibial marker; M = lateral femoral marker.

tion in the 90 degrees flexed knee.32-35 Thus, LAD knee testing position lends the PCL to inhibiting lateral tibial movement on the femur15,16,18,36 when the more vertically oriented intact PCL and intercondylar eminence encounter the lateral femoral condyle’s medial edge.37 The LAD testing force is applied from medial and slightly posterior (known as medial-posterior) to lateral and slightly anterior (known as lateral-anterior) in a direction towards the anterolateral tibial tubercle, or Gerdy’s Tubercle28 (Figure 3). In response, the cadaveric knee was flexed to 90° with the femur stabilized on the aforementioned 45° inclined wedge. The tibia was posi-

Figure 2. Finish position of the PSS; T = Tibia; F = Femur; L = Line coursing parallel to tibial diaphysis through lateral tibial marker; M = lateral femoral marker.

Figure 3. The LAD test performance in the clinical setting; * = medial arm pushing proximal tibia in lateral-anterior direction towards Gerdy´s Tubercle; = lateral hand stabilizing the femur in a medial-posterior direction.

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ous in-vivo LAD pilot testing amongst eight highly experienced testers. The LAD test trial was repeated three times. Digital images of start and finish positions were captured for each trial.

Figure 4. Start position of the LAD test with PCL dissected; T = Tibia; F = Femur; L = Line coursing parallel to tibial diaphysis through anterior-medial tibial markers; M = anterior-medial femoral marker.

After performing PSS and LAD trials with an intact PCL, an experienced anatomist examined the internal condition of the cadaveric knee through arthroscopy with the specimen positioned as previously during the LAD test. Herein the anatomist confirmed the intact status of the cruciate ligaments and posterior capsule. If the anatomist had any doubt regarding the integrity of the cruciate ligaments and/or posterior capsule, then the specimen would have been excluded from the study. However, this did not occur in any specimen. Next, the PCL was completely transected across the ligament mid-substance using an internal cutting tool guided through the arthroscope. The anatomist then re-examined the internal knee compartment to ensure that the ACL and posterior capsule remained intact. Following isolated PCL transection, three LAD test trials were performed in the same fashion described above and digital images were captured at each start and finish position. The leg was then repositioned for the PSS as previously described and three trials were performed, capturing digital images as before.

Figure 5. Finish position of the LAD test with PCL dissected; T = Tibia; F = Femur; L = Line coursing parallel to tibial diaphysis through anterior-medial tibial markers; M = anterior-medial femoral marker.

tioned in neutral rotation, with the foot fixed to the treatment table using a fixation strap. For the LAD start position, the proximal tibia was then gently manually positioned in the medial-posterior direction until the medial capsular structures were taut (Figure 4). For the LAD finish position, the proximal tibia was manually translated in lateral-anterior direction towards GerdyÂ´s Tubercle (Figure 5). The proximal tibia was translated across the complete LAD test path from the start to finish positions until the first detectable manual resistance requiring a pushing force between 112 and 122 N was encountered. This force data range was based on previ-

Data Reduction and Analysis Tibial marker displacements were digitized using a custom MATLAB Program (The Mathworks, Inc, Natwick, MA USA). The custom MATLAB program prompted the user to select a baseline image representing a resting position. In this image, a section of a standardized ruler of known distance (10-20 cm) was used to calibrate the subsequent measures. Using the coordinates of the two reference points, the total length of the selected segment was then calculated using the Pythagorean theorem. The selected distance was then converted to pixels/mm in order to proceed with further measurements. Three points on this baseline image were then chosen: two fixed points on the tibia, followed by one point on the femur. Each measurement point was established by zooming in on the intersection between the exact same four representative image pixels, reducing any error to less than a fraction of one pixel size. Lines were then extrapolated from the selected points in

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order to create a 90° angle between the points. The coordinates of a fourth point were calculated as the intersection of the two lines. The distance from this fourth point to the femoral marker was then calculated. A second image representing the end translation position was then chosen and the procedure was repeated exactly as for the baseline image. Using the two calculated distances from point four to the femoral marker, the total change in distance between the two images was then calculated in millimeters. Several investigators have previously published studies using this standardized measurement procedure across different joint systems and tissues. Gilbert et al.38,39 used the same measurement procedure when examining the uniplanar translation of cadaveric lumbar nerve roots during a clinical straight leg raise procedure. Similarly, Lohman et al.40,41 used this approach to examine the uniplanar translation of cadaveric cervical nerve roots during upper extremity neural tension testing. In both cases, embedded nerve root marker translation was measured with the same process using digital fluoroscopic images. Moreover, Wilhelm et al.42 measured uniplanar cadaveric iliotibial band deformation mounted in a load cell during clinical-grade stretching, where markers were measured in the same fashion on digital images. Finally, the inter-tester reliability and construct validity of this uniplanar measurement approach has been previously established by Cobb et al.43 who applied the approach to uniplanar tarsal bone translation during different foot postures. Their investigation tested the reliability of more than one digitizer and every single image was normalized to an object of known dimension, thus establishing the validity of this uniplanar analysis process for any image capture system at any location. In each study, significant changes in uniplanar translation were consistently observed in response to the respective study’s change in conditions. Data analyses were completed using the SPSS 22 (IBM Corp, Armonk NY, USA) software. Descriptive data, being the mean of the three trials from each test (PSS and LAD), were established for each condition. A Shapiro-Wilk test was used to establish data normality. Because the pre- versus postcut test data violated the assumption of normality,

a within-specimen Wilcoxon Signed-Rank test was used to test for significant differences in tibial translation between the intact and cut PCL conditions. The translation differences data were normally distributed, so a within-specimen t-Test was utilized to test for differences between the LAD versus PSS test conditions. The PSS values were used as a reference standard. Due to potential errors in measurement accuracy, all values were rounded to one decimal figure. Significance was set at α ≤ 0.05. RESULTS Specimen A total of 18 cadaveric specimens were examined and no specimen was eliminated in response to exclusion criteria. Digital Images A total of 24 digital images were captured per specimen [3 trials x 2 positions (start position vs. end position) x 2 PCL conditions (intact vs. cut) x 2 tests (PSS vs. LAD)]. All in all, 432 images were analyzed via MATLAB as above. LAD Testing Force Analysis The mean (SD, 95%CI) pushing force while performing the LAD test in the PCL-Intact condition was 115.5 N (1.6; 114.6 – 116.2). The mean (SD, 95%CI) pushing force while performing the LAD test with PCL-Cut was 115.1 N (2.3; 114 – 116.2). There was no statistically significant difference in LAD force values between the PCL-Intact versus PCL-Cut conditions (u = - 0.259; p = 0.795; Figure 6)

Figure 6. Testing Power for LAD test; PCL = Posterior Cruciate Ligament; Cut = Dissected; N = Newton; Error Bars = 95% Conﬁdence Interval.

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Table 1. Descriptive Data; LAD = Lateral-Anterior Drawer Test; PSS = Posterior Sag Sign; PCL = Posterior Cruciate Ligament; Cut = Dissected; SD = Standard Deviation; 95%CI = 95% Conﬁdence Interval

Lateral-anterior drawer test Proximal tibial lateral-anterior translation produced during the LAD (LADtrans) during the PCL-Intact condition was compared to the same during the PCL-Cut condition. The LADtrans during the PCLCut condition was significantly greater (u = -3.680, p = 0.002) than the LADtrans during the PCL-Intact condition (Table 1). Posterior sag sign Proximal tibial anterior-posterior translation produced during the PSS (PSStrans) during the PCLIntact condition was compared to the same during the PCL-Cut condition. The PSStrans during the PCL-Cut condition was significantly greater (u = -3.724, p = 0.002) than the PSStrans during the PCLIntact condition (Table 1). Concurrence of the LAD and the PSS In order to compare the LAD test outcomes to the outcomes of a widely accepted clinical PCL test, PSS values were used as a reference standard.22 Based on t-Test results, the tibial translation difference between PCL-Intact and PCL-Cut during the LAD was not significantly different (t = 2.029, p = 0.07) in magnitude from the same tibial translation difference during the PSS (Table 1). In order to rule out a Type II error, a post-hoc power analysis (Cohen´s r) confirmed an effect size of r = 0.3. DISCUSSION The statistically significant increase in lateral-anterior tibial translation during the LAD test after isolated PCL transection supports the hypothesis that there is an objective causality between PCL integrity and amount of tibial translation detected by the LAD

Figure 7. Difference in mean tibial translation PCL-Intact versus PCL-Cut; Diff = Difference; mm = Millimeter; Cut = Dissected; PCL = Posterior Cruciate Ligament; Error Bars = 95% Conﬁdence Interval.

test. The mean increase of 2.9mm in lateral-anterior translation represents a 38% increase in this lateralanterior movement. Thus, the results of the current study show that the LAD test can detect PCL deficient knees in this in-vitro model. Other in-vivo investigations support the current findings.15,16 Li et al.15 and van de Velde et al.16 biomechanically demonstrated that the PCL exhibits an important control function in the medial-lateral direction. In 90 degrees of knee flexion the amount of lateral tibial translation increased approximately 1.1mm15 and 1.9mm16, respectively, in patients with PCL-deficient knees when performing a single-leg lunge. While these published lateral translational values are not large, they may be relevant in the knee. The knee exhibits such tight congruencies in the collateral directions throughout the flexion-extension range. Therefore, any increase in lateral translation outside of the normal condition could be troublesome.

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A possible explanation for the higher translation values witnessed in the lateral-anterior direction during the PCL-Cut condition in the current study could be that Li et al.15 and van de Velde et al.16 conducted their tests in-vivo. This is in contrast with cadaveric testing, where no muscular knee protection or dynamic stabilization occurs and the only stability is provided by passive structures (such as patellar retinacula). Thus, the subjects’ knees in Li et al.15 and van de Velde et al.16 that were actively flexed during single leg lunges were potentially constrained through active and passive structures, lending to knee joint centering and decreased joint translations.44 Moreover, these investigators only examined pure lateral tibial translation versus the lateral-anterior test in the current investigation. In the current study, the translation force was directed in an lateral-anterior direction in order to bypass the intercondylar eminence, which can serve as a structural constraint to lateral translation within the knee joint37 and may partially explain decreased translation found in the aforementioned studies. Considering the increase in lateral translation after PCL transection that was found in the current study, the LAD test could provide the clinician with a clinical appreciation for 3-D biomechanical compromise in response to PCL-rupture. While previously reported translational values are small, this presence of lateral-anterior compromise speaks to the threedimensional constraining capability of the PCL and the LAD test could provide the clinician with clinical insight into the knee’s three-dimensional stability failure in response to a PCL injury. In the current study, the posterior tibial translation during the PSS significantly increased after the PCL was cut, which represents a 72% displacement increase. Because the PSS is commonly judged dichotomously, objective data regarding the effect of isolated PCL transection on posterior tibial displacement during the PSS is limited, creating a challenge for making any direct comparison between the current data and other investigators’ findings. In response, the authors must contrast the current PSS data with posterior drawer test data from other studies, where documented posterior tibial translation is well objectified. The current data amplitudes are found to be slightly smaller compared to other

posterior drawer test displacement data45-47 but any small translation differences could be explained by dissimilar test characteristics. In contrast to the posterior drawer test biomechanical and manual force application strategies of Bergfeld et al.45 and Fontbote et al.47, the present study did not apply any posteriorly directed force onto the tibia. Rather, by definition, only gravity created the free-floating posterior tibial translation during the PSS procedure. As one would expect, posterior tibial translation was less during the PSS versus previously documented posterior drawer test findings, where a posteriorly directed force was applied to the tibia.45,47 The current findings suggest that the LAD and PSS discover similar tibial translation changes in response to isolated PCL transection, as the LAD translation data were not significantly different (1.7mm) from the PSS data. The PSS was chosen as a comparison because it is well recognized,22 and while the quadriceps active test could serve as a comparison, it was not applicable to our specimen sample because it requires an in-vivo quadriceps activation that is unavailable in cadaveric specimens.12,14 Similarly, while the posterior drawer test could be used for comparison, the test results are difficult to judge, due to a poorly recognizable starting position for the posterior drawer test.48 Moreover, the PSS fulfills all the criteria required of a meaningful comparison test.49 The PSS is independent from the LAD test, meaning that investigators did not use the PSS results to determine LAD test results. Additionally, since in the PSS tibial translation is judged in straight posterior direction, the PSS does not include the LAD testing direction. Furthermore, the PSS exhibits a consistency with the examined LAD test concerning the structure to be evaluated and it presents with a standardized performance procedure. Moreover, investigators have reported respectable sensitivity (79%) and specificity (100%) for the PSS.22,48,50 Finally, post-hoc effect size calculations minimized any suspicion for a Type II error, suggesting a sufficient sample size and trustable findings of non-significance. Investigators have reported the value of the PSS for clinical testing PCL failures in certain circumstances. However, the PSS could be more difficult to judge when the patient has a history of Osgood-Schlatter´s

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disease, which is the most common apophysitis in pre-pubescent athletes with a prevalence of 21% in young athletes.51 Moreover, if there is swelling or obesity, the landmarks may be more difficult to locate. In order to reliably interpret the amount of any posterior tibial sag in the PCL deficient knee, the examiner has to be able to visually measure the anterior tibial plateau as the referred landmark for PSS judgment. The bony prominence in patients presenting with a history of Osgood-Schlatter´s disease would possibly reduce any visually measurable displacement of tibia on femur in the sagittal plane. Diagnosis of PCL injuries can be difficult, therefore, additional clinical tests are needed.23,52 Correspondingly, an in-vivo study for assessing inter-rater and intra-rater reliability, as well as the clinical accuracy and criterion validity of the LAD test with respect to MRI findings, is forthcoming. In order to increase the clinician’s confidence in clinical detecting PCL disruption, the LAD could be added to the toolbox and clustered with the PSS for detecting PCL tears with greater confidence. This is complemented by the practicality of the LAD test, where the manual contact of the test provides the examiner with an appreciation of the tissue integrity associated with the tibial translation. These features can be useful to aid clinician’s confidence in making the diagnosis, especially considering the lower incidence of this trauma over ACL injury.53 Because of the PCL’s lower injury rate, most clinicians do not have the opportunity to witness this injury on a frequent basis. Thus, more testing options could help them trust their judgment,52 which could lead to faster and better utilization of more confirming test procedures, such as MRI. The value of accurate and timely PCL failure detection has been debated. Patients with isolated PCL tears appear to perform at quite high levels and are less often treated surgically in the acute phase.54,55 Yet, other afflictions may arise over time as a consequence of missed isolated PCL tears, such as cartilage and meniscal degeneration. Hamada et al.56 evaluated 61 patients with isolated PCL deficient knees, witnessing extensive cartilage and meniscal degeneration as early as three months after the initial PCL rupture. Geissler and Whipple57 investigated acute and chronic isolated PCL deficient knees, reporting meniscal lesions in 36% and cartilage degeneration

in 49% of their patient population presenting with chronic isolated PCL ruptures. Moreover, van de Velde et al.16 observed increased cartilage deformation during single leg lunges in patients with PCL ruptures using MRI and dual orthogonal fluoroscopic imaging. These findings add to the value of early detection of PCL deficiency and the LAD may serve to enhance the clinician’s detection of that condition in order to allow for clinical decision making regarding interventions. LIMITATIONS AND FUTURE RESEARCH The study presents with several limitations. First, the LAD versus PSS testing order was not randomized. While this could have influenced the test outcomes, randomizing the trials would have required more frequent camera angle changes, which could have increased the risk of data collection error. Thus, the authors chose to exercise the consistent data collection in the order that was previously documented. The second limitation to this study is found in a lack of absolute blinding, where the examiner was not completely blinded to the test outcomes. However, open eyes and no complete blinding were used because the examiner was required to observe the hand placement and limb alignment throughout the test in order to maximize test consistency. Additionally, based on the location of the markers and the orientation of the camera, the examiner was not able to visualize the marker translation in the respective movement planes for each test. Moreover, while blinding was not achieved during the test procedures, image digitization was conducted by an independent investigator at a different location and time point, without knowledge of specimens in attempt to reduce any measurement bias. Third, when interpreting the results of a cadaver study, it is important to note that the role of muscle tension and/or joint forces has not been included. Thus, any application to an in-vivo situation is premature. However, the major goal of this investigation was to examine if the proposed LAD test was able to demonstrate an increased tibial translation in lateral-anterior direction in a PCL-insufficient knee. The LAD test shows promising results, but before it could be recommended for routine use, it should be

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further evaluated. Future cadaveric research could focus on a comparison of the LAD to the posterior drawer test and include accurate biomechanical testing with more precision technology, such as a material testing system (MTS). Future in-vivo studies should establish both the inter-rater and intrarater reliability of the LAD test with subjects who have sustained a PCL-rupture. Because the influence of other ligament injures on LAD response has not been tested, future research should include a measure of other ligament failures (such as ACL) on lateral-anterior tibial displacement to further elucidate the role of PCL failure in test outcomes. Furthermore, future research should include an in-vivo comparison of the LAD test versus PSS and posterior drawer test for diagnosing PCL-ruptures. Novice versus expert testers using the LAD test in-vivo should be compared in terms of their diagnostic decisionmaking. Since the meniscofemoral ligaments, if present, are assumed to be a secondary stabilizer to at least posterior tibial translation,58 future research should investigate the effect of meniscofemoral ligaments on lateral knee stability and their influence on LAD test outcomes in the PCL-deficient knee. CONCLUSION In summary, the LAD and PSS both detected significantly increased laxity in this in-vitro model that examined isolated PCL disruption. These findings are complemented by the practicality of the LAD test, where the manual contact of the test provides the examiner with an appreciation of the tissue integrity associated with the tibial translation. Additionally, the LAD in this in-vitro model suggests an appreciation of 3-D biomechanical compromise in response to isolated PCL-rupture. Moreover, the LAD may provide the examiner additional test information after PCL injury when severe knee joint swelling or increased quadriceps muscle tone may hinder accurate PSS or posterior drawer test interpretation of the extent of abnormal tibial translation. Furthermore, the tibial step-off that must be visualized during the PSS and/or palpated during the posterior drawer test may be diminished in the presence of obesity, swelling, increased muscle tone or tibial tubercle hyperostosis, making a clinical interpretation of those tests difficult for the novice examiner. The results of the current study offer biomechanical

support to carrying out future in-vivo studies that examine the utility of the LAD in clinically diagnosing PCL injuries. Once fully investigated, the LAD test could be clustered with other tests to assist in a clinician’s examination of a patient with suspected PCL-rupture. REFERENCES 1. Krüger-Franke M. Das Kniegelenk. In: Engelhardt M, ed. Sportverletzungen - Diagnose, Management und Begleitmaßnahmen. Vol 2. überarbeitete Auﬂage. München: Elsevier GmbH, Urban & Fischer Verlag; 2009:285. 2. Gläser H, Henke T. Sportunfälle - Häuﬁgkeit, Kosten, Prävention Düsseldorf 2001. 3. Strobel MJ, Weiler A, Eichhorn HJ. Diagnostik und Therapie der frischen und chronischen hinteren Kreuzbandläsion. Chirurg. 2000;71:1066-1081. 4. Schulz MS, Russe K, Weiler A, Eichhorn HJ, Strobel MJ. Epidemiology of posterior cruciate ligament injuries. Arch Orthop Trauma Surg. 2003;123(4):186-191. 5. Ruße K, Schulz MS, Strobel MJ. Epidemiologie der hinteren Kreuzbandverletzung. Arthroskopie. 2006;19(3):215-220. 6. Fanelli GC. Posterior Cruciate Ligament Injuries in Trauma Patients. Arthroscopy. 1993;9(3):291-294. 7. Fanelli GC, Edson CJ. Posterior Cruciate Ligament Injuries in Trauma Patients: Part 2. Arthroscopy. 1995;11(5):526-529. 8. Fanelli G, Beck J, Edson C. Current Concepts Review: The Posterior Cruciate Ligament. J Knee Surg. 2010;23(02):061-072. 9. Bollen S. Epidemiology of knee injuries: diagnosis and triage. Br J Sports Med. 2000;34:227-228. 10. Shelbourne KD, Davis TJ, Patel DV. The natural history of acute, isolated, nonoperatively treated posterior cruciate ligament injuries. A prospective study. Am J Sports Med. 1999;27(3):276-283. 11. Schmickal T, von Recum J, Hochstein P. Anatomie, Biomechanik und Therapie der Verletzungen des hinteren Kreuzbands. Trauma Berufskrankh. 2002;4(1):13-19. 12. Lopez-Vidriero E, Simon DA, Johnson DH. Initial Evaluation of Posterior Cruciate Ligament Injuries: History, Physical Examination, Imaging Studies, surgical and Nonsurgical Indications. Sports Med Arthrosc Rev. 2010;18:230-237. 13. Jung TM, Strobel MJ, Weiler A. Diagnostics and treatment of posterior cruciate ligament injuries. Unfallchirurg. 2006;109(1):41-59.

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14. Rigby J, Porter K. Posterior cruciate ligament injuries. Trauma. 2010;12(3):175-181. 15. Li G, Papannagari R, Li M, et al. Effect of posterior cruciate ligament deﬁciency on in vivo translation and rotation of the knee during weightbearing ﬂexion. Am J Sports Med. 2008;36(3):474-479.

27. Ulmer M, Rose T, Imhoff A. Bandverletzungen am Kniegelenk. Orthopädie und Unfallchirurgie up2date. 2006;1(5):391-410. 28. Winkel D. Orthopedische geneeskunde en manuele therapie. Deel 2 Diagnostiek extremiteiten. Houten/ Zaventem: Bohn Staﬂeu van Loghum; 1992.

16. Van de Velde SK, Bingham JT, Gill TJ, Li G. Analysis of tibiofemoral cartilage deformation in the posterior cruciate ligament-deﬁcient knee. J Bone Joint Surg. 2009;91-A(1):167-175.

29. Benkhadra M, Bouchot A, Gerard J, et al. Flexibility of Thiel’s embalmed cadavers: the explanation is probably in the muscles. Surgical and radiologic anatomy : SRA. 2011;33(4):365-368.

17. Voos JE, Mauro CS, Wente T, Warren RF, Wickiewicz TL. Posterior cruciate ligament: anatomy, biomechanics, and outcomes. Am J Sports Med. 2012;40(1):222-231.

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18. Miyasaka T, Matsumoto H, Suda Y, Otani T, Toyama Y. Coordination of the anterior and posterior cruciate ligaments in contstraining the varus-valgus and internal-external rotatory instability of the knee. J Orthop Sci. 2002;7:348-353. 19. Ochi M, Murao T, Sumen Y, Kobayashi K, Adachi N. Isolated Posterior Cruciate Ligament Insufficiency Induces Morphological Changes of Anterior Cruciate Ligament Collagen Fibrils. Arthroscopy. 1999;15(3):292-296. 20. Butler DL, Noyes FR, Grood ES. Ligamentous restraints to anterior-posterior drawer in the human knee. A biomechanical study. J Bone Joint Surg. 1980;62-A:259-270. 21. LaPrade CM, Civitarese DM, Rasmussen MT, LaPrade RF. Emerging Updates on the Posterior Cruciate Ligament: A Review of the Current Literature. Am J Sports Med. 2015; 43(12):3077-3092. 22. Malanga GA, Andrus S, Nadler SF, McLean J. Physical examination of the knee: a review of the original test description and scientific validity of common orthopedic tests. Arch Phys Med Rehabil. 2003;84(4):592-603. 23. Kopkow C, Freiberg A, Kirschner S, Seidler A, Schmitt J. Physical Examination Tests for the Diagnosis of Posterior Cruciate Ligament Rupture: A Systematic Review. J Orthop Sports Phys Ther. 2013;43(11):804-814. 24. Margheritini F, Rhin J, Musahl V, Mariani PP, Harner C. Posterior Cruciate Ligament Injuries in the Athlete. An Anatomical, Biomechanical and Clinical Review. Sports Med. 2002;32(6):393-408. 25. Pierce CM, O’Brien L, Griffin LW, Laprade RF. Posterior cruciate ligament tears: functional and postoperative rehabilitation. Knee Surg Sports Traumatol Arthrosc. 2013;21(5):1071-1084. 26. Cooper DE. Clinical Evaluation of Posterior Cruciate Ligament Injuries. Sports Med Arthrosc Rev. 1999;7:243-252.

31. Streeck U, Focke J, Klimpel L, Noack DW. Manuelle Therapie und komplexe Rehabilitation. Band 2: Untere Körperregion. In: Streeck U, Focke J, Klimpel L, Noack DW, 1st ed. Heidelberg: Springer Medizin Verlag; 2007. 32. Kapandji IA. Funktionelle Anatomie der Gelenke. Band 2: Untere Extremität. 3., unveränderte Auﬂage. Stuttgart: Hippokrates Verlag; 2001. 33. Blankevoort L, Huiskes R, de Lange A. Recruitment of Knee Joint Ligaments. J Biomech Eng. 1991;113:94103. 34. Williams A, Logan M. Understanding tibio-femoral motion. Knee. 2004;11(2):81-88. 35. Amis AA, Gupte CM, Bull AM, Edwards A. Anatomy of the posterior cruciate ligament and the meniscofemoral ligaments. Knee Surg Sports Traumatol Arthrosc. 2006;14(3):257-263. 36. Cyriax J. Textbook of Orthopaedic Medicine. Volume One: Diagnostic of Soft Tissue Lesions. 7th Edition. London: Ballière Tindall; 1978. 37. Goodfellow J, O´Connor J. The Mechanics of the Knee and Prothesis Design. J Bone Joint Surg. 1978;60-Br:358-369. 38. Gilbert KK, Brismée J-M, Collins DL, et al. 2006 Young Investigator Award Winner: Lumbosacral Nerve Root Displacement and Strain. Part 1. A Novel Measurement Technique During Straight Leg Raise in Unembalmed Cadavers. Spine (Phila Pa 1976). 2007;32(14):1513-1520. 39. Gilbert KK, Brismée J-M, Collins DL, et al. 2006 Young Investigator Award Winner: Lumbosacral Nerve Root Displacement and Strain. Part 2. A Comparison of 2 Straight Leg Raise Conditions in Unembalmed Cadavers. Spine (Phila Pa 1976). 2007;32(14):1521-1525. 40. Lohman CM, Gilbert KK, Sobczak S, et al. 2015 Young Investigator Award Winner: Cervical Nerve Root Displacement and Strain During Upper Limb Neural Tension Testing: Part 1: A Minimally Invasive

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Assessment in Unembalmed Cadavers. Spine (Phila Pa 1976). 2015;40(11):793-800. Lohman CM, Gilbert KK, Sobczak S, et al. 2015 Young Investigator Award Winner: Cervical Nerve Root Displacement and Strain During Upper Limb Neural Tension Testing: Part 2: Role of Foraminal Ligaments in the Cervical Spine. Spine (Phila Pa 1976). 2015;40(11):801-808. Wilhelm M, Matthijs O, Browne K, et al. Deformation Response of the Iliotibial Band-Tensor Fascia Lata Complex to Clinical-Grade Longitudinal Tension Loading In-Vitro. Int J Sports Physical Ther. 2017;12(1):16-24. Cobb SC, James CR, Hjertstedt M, Kruk J. A Digital Photographic Measurement Method for Quantifying Foot Posture: Validity, Reliability, and Descriptive Data. J Athl Train. 2011;46(1):20-30. Kibler WB, Livingston B. Closed-chain rehabilitation for upper and lower extremities. J Am Acad Orthop Surg. 2001;9:412-421. Bergfeld JA, McAllister DR, Parker RD, Valdevit ADC, Kambic H. The effects of tibial rotation on posterior translation in knees in which the posterior cruciate ligament has been cut. J Bone Joint Surg. 2001;83-A(9):1339-1343. Logan M. The Effect of Posterior Cruciate Ligament Deﬁciency on Knee Kinematics. Am J Sports Med. 2004;32(8):1915-1922. Fontbote CA, Sell TC, Laudner KG, et al. Neuromuscular and biomechanical adaptations of patients with isolated deﬁciency of the posterior cruciate ligament. Am J Sports Med. 2005;33(7):982989. Rubinstein RA, Shelbourne KD, McCarroll JR, VanMeter CD, Rettig AC. The Accuracy of the Clinical Examination in the Setting of Posterior Cruciate Ligament Injuries. Am J Sports Med. 1994;22(4):550-557.

49. Fritz JM, Wainner RS. Examining Diagnostic Tests: An Evidence-Based Perspective. Phys Ther. 2001;81:1546-1564. 50. Fowler PJ, Messieh SS. Isolated posterior cruciate ligament injuries in athletes. Am J Sports Med. 1987;15(6):553-557. 51. Kujala UM, Kvist M, Heinonen O. Osgood-Schlatter´s disease in adolescent athletes. Retrospective study of incidence and duration. Am J Sports Med. 1985;13(4):236-241. 52. Moulton SG, Cram TR, James EW, Dornan GJ, Kennedy NI, LaPrade RF. The Supine Internal Rotation Test: A Pilot Study Evaluating Tibial Internal Rotation in Grade III Posterior Cruciate Ligament Tears. J Orthop Sports Med. 2015;3(2). 53. Colvin AC, Meislin RJ. Posterior Cruciate Ligament Injuries in the Athlete; Diagnosis and Treatment. Bull NYU Hosp Jt Dis. 2009;67(1):45-51. 54. Patel DV, Allen AA, Warren RF, Wickiewicz TL, Simonian PT. The nonoperative treatment of acute, isolated (partial or complete) posterior cruciate ligament-deﬁcient knees: an intermediate-term follow-up study. HSS J. 2007;3(2):137-146. 55. Torg JS, Barton TM, Pavlov H, Stine R. Natural History of the Posterior Cruciate Ligament-Deﬁcient Knee. Clin Orthop Relat Res. 1989;246:208-216. 56. Hamada M, Shino K, Mitsuoka T, Toritsuka Y, Natsu-ume T, Horibe S. Chondral Injury Associated with Acute Isolated Posterior Cruciate Ligament Injury. Atrhroscopy. 2000;16(1):59-63. 57. Geissler WE, Whipple TL. Intraarticular abnormalities in association with posterior cruciate ligament injuries. Am J Sports Med. 1993;21(6):846-849. 58. Gupte CM, Bull AMJ, Thomas RD, Amis AA. The meniscofemoral ligaments: secondary restraints to the posterior drawer. J Bone Joint Surg. 2003;85-B:765773.

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IJSPT

ORIGINAL RESEARCH

RELATIONSHIPS AMONG COMMON VISION AND VESTIBULAR TESTS IN HEALTHY RECREATIONAL ATHLETES John D. Heick, PT, PhD, DPT, OCS, NCS, SCS1 Curt Bay, PhD2 Thomas P. Dompier, PhD, ATC3 Tamara C. Valovich McLeod, PhD, ATC, FNATA2

ABSTRACT Background: Disruption of the visual and vestibular systems is commonly observed following concussion. Researchers have explored the utility of screening tools to identify deficits in these systems in concussed patients, but it is unclear if these tests are measuring similar or distinct phenomena. Purpose: To determine the relationships between common vestibular tests including the King-Devick (K-D) test, Sensory Organization Test (SOT), Head Shake-Sensory Organization Test (HS-SOT), and Dynamic Visual Acuity (DVA) test, when administered contiguously, to healthy recreational athletes aged 14 to 24 years. Study Design: This study used a prospective design to evaluate relationships between the K-D, SOT, HSSOT, and DVA tests in 60 healthy individuals. Methods: Sixty participants (30 males, 30 females; mean age, 19.9Âą3.74 years) completed the four tests in a single testing session. Results: Results did not support a relationship between any pair of the K-D, SOT, HS-SOT, and DVA tests. Pearson correlations between tests were poor, ranging from 0.14 to 0.20. As expected the relationship between condition 2 of the SOT and HS-SOT fixed was strong (ICC=0.81) as well as condition 5 of the SOT with HS-SOT sway (ICC=0.78). The test-retest reliability of all 4 tests was evaluated to ensure the relationships of the 4 tests were consistent between test trials and reliability was excellent with intraclass correlations ranging from 0.79 to 0.97. Conclusions: The lack of relationships in these tests is clinically important because it suggests that the tests evaluate different aspects of visual and vestibular function. Further, these results suggest that a comprehensive assessment of visual and vestibular deficits following concussion may require a multifaceted approach. Level of Evidence: 2b: Individual Cohort Study. Key words: Dynamic visual acuity, saccades, vestibulo-ocular reflex

Northern Arizona University, Flagstaff, AZ, USA A.T. Still University, Mesa, AZ, USA 3 Datalys Center for Sports Injury Research and Prevention, Inc, Indianapolis, IN, USA 2

CORRESPONDING AUTHOR John D. Heick, PT, PhD, DPT, OCS, NCS, SCS Associate Professor, Physical Therapy Northern Arizona University, PO Box 15105 Flagstaff, Arizona 86011 Phone: 928-523-8394, Fax 928-523-9289 Eâ&#x20AC;&#x201C;mail: John.Heick@nau.edu

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INTRODUCTION The functional damage to the brain parenchyma from concussion commonly manifest as deficits in several domains including cognitive, postural stability, autonomic dysfunction, and behavioral. Evaluation of patients with a suspected concussion should address all these domains. The most frequently documented complaint following concussion is headache; but dizziness, postural instability, and balance dysfunction are also common.1-5 Although researchers have assessed the efficacy of screening tools to detect postural instability such as the Sensory Organization Test,6-13 the contributions of the visual and vestibular systems to postural instability in the concussed athlete have not been thoroughly investigated and are frequently studied independently of other clinical measures. Visual system disruption is frequently observed following concussion, and 65% to 90% of patients show oculomotor dysfunction,14 which may impair postural stability.6,7 Visual and vestibular system disorders have both been implicated as causes of postural instability.1,5,10 An easily administered screening tool, the King Devick (K-D) test, assesses both the visual and vestibular systems.6,7,15,16 This test also incorporates the cognitive domain because it identifies impairment of eye movements, attention, language, and other symptoms that are associated with suboptimal brain function.6,7,15,16 Researchers have suggested that the K-D test may be useful in the acute assessment of concussion.6,7,15,17,18 Coordination of the three peripheral sensory systems (visual, vestibular, and somatosensory) is required to function normally and maintain postural stability. Concussion disrupts the integration of postural responses from these systems and results in postural instability.19-22 Several studies evaluated the usefulness of postural stability tests in concussed athletes.19-21,23-28 The Sensory Organization Test (SOT) is commonly used to assess postural stability.28 The Head Shake-Sensory Organization Test (HS-SOT) modifies two of the standard SOT conditions by including dynamic head motions that stimulate the semicircular canals within the vestibular system. The Dynamic Visual Acuity (DVA) test is used to assess different aspects of vestibular function including the vestibulo-ocular reflex.29-31 Although there are multiple means of assessing vestibular and oculomotor

function, it is unclear whether these tests are assessing the same or different components of vestibular and oculomotor function. In a recent report, results from the DVA differed from those of other oculomotor tests, such as gaze stability tests, suggesting that these tests assess different constructs.29 Each of these tests (K-D, SOT, HS-SOT, and DVA) evaluates components of vestibular and oculomotor function but the degree to which they measure common deficits (or attributes) is not known. The purpose of the current study was to assess the relationships between common vestibular tests, including the K-D, SOT, HSSOT, and DVA tests, when administered contiguously, to healthy recreational athletes aged 14 to 24 years. Additionally, a secondary purpose was to conduct a test-retest reliability evaluation of the four tests. METHODS Statistical Methods This study used a prospective design to evaluate relationships between the K-D, SOT, HS-SOT, and DVA tests in 60 healthy individuals. The reliability of all four tests was evaluated to ensure the relationships of the four tests were consistent across individual test trials. Participants Participants were recruited through flyers distributed to public and private school systems in a metropolitan community between May and September 2014. Participants were required to be 14 to 24 years old and possess sufficient English language skills to complete questionnaires. Exclusion criteria were lower extremity musculoskeletal injuries in the prior three months; history of a head injury in the past year; or diagnosis of a visual, vestibular, or balance disorder. Telephone screening was used to ascertain eligibility. Participants meeting study criteria were provided information about the purpose of the research and the potential risks. Participants provided written informed consent or the parent/guardian provided written permission and the child assented. Experimental procedures were approved by the institutional review boards associated with the study. Participants completed a personal/medical history form prior to testing to ensure there were no reasons for exclusion (Appendix 1).

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INSTRUMENTATION King-Devick test The K-D test is a screening tool that identifies impairments of eye movements, attention, and language6,7,10 through measurement of the speed of rapid number naming. 6,7,15,16 Participants read aloud a series of single-digit numbers from left to right on one practice card and three test cards.6,7,15,16 Standardized instructions were followed. The sum of the time to complete the three test cards constitutes the summary score, known as the K-D composite score. The number of errors is recorded as the K-D error score. Sensory Organization Test The SOT was performed using the NeuroCom Equitest BalanceMaster. The SOT uses a computerized system with a servo-controlled dual force plate and visual surround to determine whether an individual can effectively use inputs from visual, somatosensory, and vestibular systems to maintain balance while suppressing inaccurate sensory information.20,25 The SOT evaluates sensory interactions during six conditions that alter visual, somatosensory, or vestibular systems with the participant attempting to maintain steady state standing balance while wearing a harness within the NeuroCom Equitest BalanceMaster. The conditions are the following: (1) eyes open standing on a firm surface; (2) eyes closed while standing on a firm surface; (3) sway-referenced vision standing on a firm surface; (4) eyes open standing on a sway-referenced surface; (5) eyes closed standing on a sway-referenced surface; and (6) eyes open, swayreferenced vision standing on a sway-referenced surface. Three trials are performed for each condition to generate a composite equilibrium score. The SOT has good-to-moderate test-retest reliability and has assessed sensory contributions to balance control in children, young adults, older adults, and individuals with neurological disorders.24,28,32,33 It has also been used to evaluate the effectiveness of interventions targeted at improving balance.26,33,34 Head Shake Sensory-Organization Test The HS-SOT was performed using the NeuroCom Equitest BalanceMaster. The HS-SOT is an enhancement of the SOT and was developed to improve the delineation of balance performance.35-38 In the HSSOT, dynamic head movements are incorporated into

standard SOT condition 2 and condition 5.35,36 Unlike the SOT where the head is static, the HS-SOT requires active head movements in the horizontal plane, as if saying no repeatedly, to correspond with visual and auditory feedback while maintaining a fixed head velocity at approximately 100Â° per second as measured by an accelerometer. In addition to assessing the possible influence of head movements on postural stability, the HS-SOT also stimulates the semicircular canals.35 This stimulation creates additional vestibular input that must be integrated during the balance task.37 Therefore, the HS-SOT may quantify subtle balance deficits and enhance the clinical standard use of the SOT. 36 The HS-SOT has been shown to have excellent test-retest reliability in healthy, younger adults and moderate-to-good test-retest reliability in healthy, older adults.39 At least five trials of the HS-SOT are required to calculate composite fixed (condition 2) and sway (condition 5) scores. Dynamic Visual Acuity test The DVA test was performed using the NeuroCom Equitest BalanceMaster. The DVA test assesses the ability to use the peripheral vestibular system for appropriate visual target capture during head movements40 and quantifies decrements in visual acuity during dynamic head tasks.30,41 In the computerized DVA test, individuals report the orientation of an optotype E (which way the open legs of the E point) displayed on the computer while their head is in motion at a fixed velocity of approximately 100Â° per second.30,31 The DVA test varies the size of the E to identify the smallest E that can be accurately recognized at least 60% of the time when the head is in motion. 41 The DVA software uses logMAR units (log of the minimum angle of resolution) to quantify the ability to accurately identify the orientation of the E.30 The computerized DVA test has been reported to detect peripheral vestibular dysfunction in adults40 and children42 and has been verified as a sensitive and specific assessment for detecting peripheral vestibular impairments.41 UCLA Dizziness Questionnaire 2 The University of California, Los Angeles Dizziness Questionnaire 2 (UCLA-DQ2) is a subjective measure of a participantâ&#x20AC;&#x2122;s dizziness.43,44 The UCLA-DQ2 uses a six-point Likert visual analogue scale that

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quantifies any symptoms associated with dizziness, from 0 (no symptoms) to 5 (severe symptoms). The Questionnaire is presented in Appendix 2.43,45 Procedures All testing was conducted at a university research laboratory and collected repeated measurements in a single visit. Participants were permitted to use glasses or contact lenses and encouraged to report any visual disorders that would affect their ability to participate. They completed a personal/medical history form and the UCLA DQ-2. Instructions were provided before each test was performed, and participants were asked to demonstrate their understanding of the test before proceeding. Participants were offered water and rest breaks between tests to ensure hydration and adequate rest. Participants were asked to rate their dizziness or any symptoms associated with dizziness using the UCLA DQ-2 between each test and during testing because of the repeated cervical motions required by the HS-SOT and DVA tests. To analyze for order effect of the four tests, two predetermined testing orders were used and participants flipped a coin to determine the order (Figure 1). The SOT and HS-SOT were considered a single test because the SOT is required to initiate the HSSOT. The K-D test was completed in a seated posi-

tion at a self-selected distance for reading, following procedures used in recent studies,6,7,16,18 for each of 3 trials (trial=1 complete test). The time required to complete each trial was measured by a stopwatch to the nearest tenth of a second, and the number of errors was recorded. The SOT was performed following a standardized procedure in the literature.46 Next, participants were returned to the BalanceMaster for the HS-SOT. Per manufacturer recommendations, participants were placed in an appropriately sized harness that did not restrict sway, and their malleoli were aligned with the axis of rotation based on their height. Participants were told to stand as steadily as possible during the six conditions. Three trials of the SOT were performed for each condition to calculate a composite score, ranging from 0 to 100. The score was more heavily weighted for the more difficult balance conditions (5 and 6), and a score of 0 was given to participants who required the harness to prevent a fall. After completing the SOT, participants were disengaged from the BalanceMaster and offered a seated rest break. Next, participants were returned to the BalanceMaster for the HS-SOT. Participants were assisted in donning the head accelerometer and instructed to move their head in the horizontal axis at a velocity of approximately 100Â° per second

Figure 1. Randomly Selected Testing Order (A or B) Used in the Study. Abbreviations: DVA= Dynamic Visual Acuity test; HSSOT= Head Shake-Sensory Organization Test; KD= King-Devick test; SOT= Sensory Organization Test. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 584

that corresponded with an auditory tone and visual feedback on the computer monitor. Participants practiced maintaining the horizontal motion at the appropriate velocity using the visual feedback until they were able to maintain the motion with their eyes closed and using the auditory feedback. Five trials of the HS-SOT were performed to calculate the composite fixed and sway scores. After completing the HS-SOT, participants were offered a seated rest break before continuing with the K-D test. For the DVA test, participants stood eight feet from the Neurocom BalanceMaster computer screen following standard procedures recommended by the manufacturer, and the screen was adjusted so that the center of the screen was at the participantâ&#x20AC;&#x2122;s eye level. Participants who wore glasses or contacts were asked to wear them. Participants were provided two practice trials of at least two minutes for the horizontal motion and the visual acuity component of the DVA. Head motion velocity and appropriate horizontal motion were indicated with visual feedback from the computer. When the participant verbalized the direction of the optotype E to the investigator, the investigator clicked a handheld mouse-like device to indicate the direction (i.e., up, down, left, or right). Time to complete the DVA test was done for three conditions: static, left, and right. The DVA acuity score represents a difference between the static DVA score and the score achieved during the dynamic components (left and right) of this test. Statistical Analyses A sample size calculation was performed to estimate modest relationships between the K-D, SOT, HS-SOT, and DVA tests using PASS version 12 software (NCSS Statistical Software, Kaysville, UT). A sample of 56 participants was used to detect an intraclass correlation coefficient (ICC) as low as 0.35 with an Îą of 0.05. Sixty participants were recruited to allow for possible dropouts. Test-retest reliabilities of the SOT composite, HS-SOT fixed composite, HS-SOT sway composite, and DVA scores were calculated by comparison of each trial within each test. The reliability of the K-D test was estimated using the K-D composite score averaged across the 3 trials. The ICCs were calculated using a 2-way, random-effects model. Reliability was defined

in terms of consistency, and the average agreement was reported. An ICC less than 0.40 indicated poor reliability, 0.40 to 0.74 indicated moderate-to-good reliability, and greater than 0.75 indicated excellent reliability.45 The strength of relationships between the tests was estimated using Pearson correlation coefficients because distributions were approximately normal. The relationship between the SOT conditions 2 and 5 and the HS-SOT fixed and sway composites was also tested. Pearson correlations less than 0.50 were considered poor, between 0.50 and 0.75 were considered moderate, and greater than 0.75 were considered good. A p-value of less than .05 (2-tailed) was considered statistically significant. SPSS version 23.0 (IBM Corp., Armonk, NY) was used to analyze the data. RESULTS Sixty healthy individuals (30 males, 30 females; mean [SD] age, 19.9 [3.74] years) participated in the current study (Table 1). There were no dropouts or excluded data. Testing order showed no effect, so was not used in the analyses. Twelve participants reported symptoms of dizziness associated with the HS-SOT. No reports exceeded a 2 of 5 (mild) rating. There were 10 reports of dizziness during the SOT and HS-SOT tests, and nine reports of dizziness during the DVA Table 1. Demographic Characteristics of the Participants (N=60) Demographic Characteristic

No. (%) or Mean (SD)a

Sex Male

30 (50)

Female

30 (50)

Age (years)

19.9 (3.74)

Glasses/contacts

23 (38.3)

Previous head injury

10 (16.6)

Sports history Caffeine intake

57 (95.0)

None

21 (35.0)

1/day

25 (41.7)

Less than 3/day

9 (15.0)

More than 3/day

5 (8.3)

Age is reported as mean (SD).

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test. Dizziness symptoms subsided within two minutes of test completion. Test-retest reliability for all tests was excellent (Table 2). Table 3 provides Pearson correlation coefficients and p-values for each pair of tests. The HS-SOT sway was significantly correlated with the SOT composite (r= -0.344). The only other correlations that exceeded r=0.25 were those calculated between subtests of the DVA, which should be strongly correlated with one another because they are measuring similar attributes. The association between condition 2 of the SOT and the HS-SOT fixed was strong (r =0.81) and the association between condition 5 and the HS-SOT sway was also strong (r=0.78). DISCUSSION Although excellent reliability was found for the K-D, HS-SOT fixed, HS-SOT sway, and DVA tests, only Table 2. Test-Retest Reliability of the King-Devick (K-D) Test, the Sensory Organization Test (SOT), Head Shake-Sensory Organization Test (HS-SOT), and the Dynamic Visual Acuity (DVA) Test in Healthy Individuals Aged 14 to 24 Years (N=60) Test

ICC (95% CI)

K-D composite score for 3 trials

0.97 (0.96 to 0.98)

SOT composite

0.83 (0.75 to 0.88)

HS-SOT fixed

0.81 (0.73 to 0.88)

HS-SOT sway

0.79 (0.69 to 0.87)

DVA static

0.83 (0.74 to 0.89)

All tests had a p<.001. Abbreviations: CI, confidence interval; ICC, intraclass correlation coefficient.

weak correlations were found between these measures in healthy recreational athletes aged 14 to 24 years. While each of the tests assessed in the current study have been studied independently as concussion assessment tools,5-9,17,19,21,40,43 results of the current study suggest multiple concussion assessment tools may need to be used together to identify all areas of suboptimal brain function after a concussion. Even though concussion assessment should address symptoms, cognition, and postural stability, many researchers and clinicians categorize postural stability tests together. Results of the current study indicate that these vestibular tests are measuring different phenomena. The results of the current study are clinically important because knowing that there are no relationships between the K-D, SOT, HS-SOT, and DVA tests may encourage healthcare professionals to appreciate the complexity of impairments to visual/vestibular function. Each of these tests seems to be measuring a specific, and seemingly independent, aspect of the visual or vestibular system. Specifically, the postural stability tests (SOT, HS-SOT, and DVA) measure the integration of distinct aspects of vision with the vestibular system, and the K-D test assesses saccadic accuracy and attention and identifies oculomotor deficits.24,47 The finding that tests seem to be measuring different attributes reinforces the notion that the K-D test should be incorporated with other visual and vestibular assessment tests as part of a multifaceted approach to identify the anatomical structures that are injured in a concussed patient and manifest as disruptions to the symptom, cognition, and postural stability domains.

Table 3. Pearson Correlation CoefďŹ cients (P values) Between the King-Devick Test (K-D), the Sensory Organization Test (SOT), the Head Shake-Sensory Organization Test (HS-SOT), and the Dynamic Visual Acuity (DVA) Test in Healthy Individuals Aged 14 to 24 Years (N=60) Test

HS-SOT Fixed

K-D HS-SOT fixed HS-SOT sway DVA static

HS-SOT Sway

DVA Static

DVA Right

DVA Left

SOT Composite

0.079 (0.550)

-0.169 (0.198)

0.117 (0.374)

0.114 (0.385)

0.243 (0.062)

0.35 (0.791)

0.084 (0.523)

0.173 (0.187)

-0.179 (0.170)

-0.116 (0.379)

-0.032 (0.807)

0.194 (0.138)

0.126 (0.339)

0.017 (0.895)

-0.344 (0.007)

0.645 (<0.001)

0.521 (<0.001)

-0.016 (0.901)

0.726 (<0.001)

-0.215 (0.099)

-0.216 (0.097)

DVA right DVA left

Data are presented as Pearson correlation coefficients with associated p-values (2-tailed).

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Concussion assessment should be performed using a multifaceted approach with measures that have sound psychometric properties. While the SOT is a recommended post-concussion assessment tool for concussed athletes,48 it is not typically used in the clinic for baseline testing, most likely due to cost of the equipment. It is primarily used to detect postural instability post-concussion to determine which of the three sensory systems (visual, vestibular, and somatosensory) is impaired.25,46,49 The HS-SOT has identified healthy young and older individuals with balance deficits.36 Pang et al36 found that the HS-SOT sway was a better indicator of postural instability than the HS-SOT fixed. The authors also noted a significant difference between HS-SOT fixed and sway compared with conditions 2 and 5 of the SOT.36 In the current study, the association between condition 2 of the SOT and the HS-SOT fixed was strong (r =0.81). Regarding reliabilities for the HS-SOT condition 2 and 5, results of the current study were slightly lower than those presented in the Pang et al 36 study for younger healthy adults (ICC=0.85) but higher than those presented for older adults (ICC=0.64). A similar association for younger adults for the HS-SOT sway (ICC=0.81) was found in the current study,36 but results were higher for younger healthy adults (ICC=0.78) and older healthy adults (ICC=0.55). These differences may be the result of fatigue. Participants in the current study performed all six conditions of the SOT (18 trials) prior to the HS-SOT as opposed to performing conditions 2 and 5 of the SOT (6 trials) prior to performing the HS-SOT. In the current study, performance of all trials of the SOT and HS-SOT was completed because that is how the tests are performed in the clinical setting. The HS-SOT and DVA tests both use high-velocity, active head movements to assess the function of the semicircular canals and the vestibulo-ocular reflex.39,50,51 These high-velocity movements may dissuade clinicians from using these tests on a concussed athlete since they may make the athlete’s symptoms worse because of the required active head motions. The DVA test uses the same active head motions as the HS-SOT but identifies peripheral vestibular

deficits by increasing retinal image slip during horizontal plane head motion.40,52-54 Retinal image slip occurs when the gain of the vestibulo-ocular reflex is decreased. Gain is analogous to controlling the volume of a radio; as the gain (volume) decreases, the vestibular-ocular reflex decreases. Therefore, the image on the retina is unstable and may manifest as blurred vision or oscillopsia. In a healthy individual without a vestibular deficit, the active head motion matches the eye motion during vestibulo-ocular reflex testing. When a vestibular lesion is present, such as in a concussion, the eye motion does not match the head motion. If retinal image slip velocity exceeds 2° to 4° per second, the visual acuity is reduced47 and can be measured by errors in DVA testing. In the current study, assessment was conducted on healthy young adults without known visual or vestibular deficits, such as retinal slip or complaints of blurred vision. Thirty-eight percent of participants wore corrective lenses for the DVA test. Although the K-D and DVA tests are not considered neurocognitive tests, both have components of cognitive measurement. Vital et al51 suggested that the DVA test should be performed passively because active horizontal plane motion relies on smooth pursuit and saccades, which may contribute to improved gaze stabilization, whereas passive horizontal plane motion is more sensitive. Future studies should investigate the use of the DVA test in a large sample of healthy young adults to determine if passive horizontal plane motion is more reliable than active horizontal plane motion. Such information would be useful before testing a concussed athlete since the active horizontal plane head motion of the DVA may increase the concussed athlete’s dizziness. The current study had several limitations. The participants were non-concussed healthy participants without vision or vestibular deficits, and participants were not required to be athletes. Therefore, it is unknown if the findings translate to concussed athletes. Further, the software used in the current study limited the number of trials that could be performed for the six conditions of the SOT. Future studies should consider extending the number of trials beyond the standard three trials for the K-D test and the SOT to determine if a performance plateau has been reached. Finally, the SOT has established normative values in healthy individuals for a variety

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of age ranges, but no age-appropriate, normative values have been established for the K-D test, the HS-SOT, or the DVA test. Determining normative values for these tests would allow clinicians to identify unusual scores at baseline and post-concussion. CONCLUSIONS In the current study, the K-D test, the SOT, the HSSOT, and the DVA test were investigated to determine if there were relationships between these tests in a sample of healthy recreational athletes aged 14 to 24 years. There were no significant relationships between the tests, which suggests that the tests are measuring different aspects of the visual and vestibular system. Based on the current study, each test measures a distinct anatomical structure’s ability to perform a specific function. It is not yet clear which tests or combination of tests are optimal to detect impairments following a concussion. REFERENCES 1. Alsalaheen BA, Mucha A, Morris LO, et al. Vestibular rehabilitation for dizziness and balance disorders after concussion. J Neurol Phys Ther. 2010;34(2): 87-93. 2. Gurr B, Moffat N. Psychological consequences of vertigo and the effectiveness of vestibular rehabilitation for brain injury patients. Brain Inj. 2001;15(5):387-400. 3. Guskiewicz KM, McCrea M, Marshall SW, et al. Cumulative effects associated with recurrent concussion in collegiate football players: the NCAA Concussion Study. JAMA. 2003;290(19):2549-2555. 4. Herdman SJ. Treatment of benign paroxysmal positional vertigo. Phys Ther. 1990;70(6):381-388. 5. Hoffer ME, Gottshall KR, Moore R, Balough BJ, Wester D. Characterizing and treating dizziness after mild head trauma. Otol Neurotol. 2004;25(2):135-138. 6. Galetta KM, Barrett J, Allen M, et al. The KingDevick test as a determinant of head trauma and concussion in boxers and MMA ﬁghters. Neurology. 2011;76(17):1456-1462. 7. Galetta KM, Brandes LE, Maki K, et al. The KingDevick test and sports-related concussion: study of a rapid visual screening tool in a collegiate cohort. J Neurol Sci. 2011;309(1-2):34-39. 8. Gottshall K, Drake A, Gray N, McDonald E, Hoffer ME. Objective vestibular tests as outcome measures in head injury patients. Laryngoscope. 2003;113(10):1746-1750.

9. McCrea M, Guskiewicz KM, Marshall SW, et al. Acute effects and recovery time following concussion in collegiate football players: the NCAA Concussion Study. JAMA. 2003;290(19):2556-2563. 10. Ventura RE, Balcer LJ, Galetta SL. The neuroophthalmology of head trauma. Lancet Neurol. 2014;13(10):1006-1016. 11. Ventura RE, Jancuska JM, Balcer LJ, Galetta SL. Diagnostic tests for concussion: is vision part of the puzzle? J Neuro-Ophthalmol. 2015;35(1):73-81. 12. Dickin DC. Obtaining reliable performance measures on the sensory organization test: altered testing sequences in young adults. Clin J Sport Med. 2010;20(4):278-285. 13. Park MK, Lim HW, Cho JG, Choi CJ, Hwang SJ, Chae SW. A head shake sensory organization test to improve the sensitivity of the sensory organization test in the elderly. Otol Neurotol. 2012;33(1):67-71. 14. Heitger MH, Anderson TJ, Jones RD, DalrympleAlford JC, Frampton CM, Ardagh MW. Eye movement and visuomotor arm movement deﬁcits following mild closed head injury. Brain. 2004;127(Pt 3):575-590. 15. King D, Clark T, Gissane C. Use of a rapid visual screening tool for the assessment of concussion in amateur rugby league: a pilot study. J Neurol Sci. 2012;320(1-2):16-21. 16. Tjarks BJ, Dorman JC, Valentine VD, et al. Comparison and utility of King-Devick and ImPACT(R) composite scores in adolescent concussion patients. J Neurol Sci. 2013;334(1-2):148-153. 17. Braun PA, Kaminski TW, Swanik CB, Knight CA. Oculomotor function in collegiate student athletes with a previous history of sport-related concussion. Athl Train Sports Health Care. 2013;6(5):282-288. 18. Vartiainen MV, Holm A, Peltonen K, Luoto TM, Iverson GL, Hokkanen L. King-Devick test normative reference values for professional male ice hockey players. Scand J Med Sci Sports. 2015;25(3):e327-330. 19. Guskiewicz KM. Postural stability assessment following concussion: one piece of the puzzle. Clin J Sport Med. 2001;11(3):182-189. 20. Guskiewicz KM. Assessment of postural stability following sport-related concussion. Curr Sports Med Rep. 2003;2(1):24-30. 21. Guskiewicz KM, Ross SE, Marshall SW. Postural Stability and Neuropsychological Deﬁcits After Concussion in Collegiate Athletes. J Athl Train. 2001;36(3):263-273. 22. Powers KC, Kalmar JM, Cinelli ME. Recovery of static stability following a concussion. Gait Posture. 2014;39(1):611-614.

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23. Evans MK, Krebs DE. Posturography does not test vestibulospinal function. Otolaryngol Head Neck Surg. 1999;120(2):164-173. 24. Ford-Smith CD, Wyman JF, Elswick RK, Jr., Fernandez T, Newton RA. Test-retest reliability of the sensory organization test in noninstitutionalized older adults. Arch Phys Med Rehabil. 1995;76(1):77-81. 25. Nashner LM, Peters JF. Dynamic posturography in the diagnosis and management of dizziness and balance disorders. Neurol Clin. 1990;8(2):331-349. 26. O’Neill DE, Gill-Body KM, Krebs DE. Posturography changes do not predict functional performance changes. Am J Otology. 1998;19(6):797-803. 27. Sinaki M, Lynn SG. Reducing the risk of falls through proprioceptive dynamic posture training in osteoporotic women with kyphotic posturing: a randomized pilot study. Am J Phys Med Rehabil. 2002;81(4):241-246. 28. Wrisley DM, Stephens MJ, Mosley S, Wojnowski A, Duffy J, Burkard R. Learning effects of repetitive administrations of the sensory organization test in healthy young adults. Arch Phys Med Rehabil. 2007;88(8):1049-1054. 29. Asken BM, Mihalik JP, Schmidt JD, Littleton AC, Guskiewicz KM, Hopfinger JB. Visual performance measures and functional implications in healthy participants: a sports concussion perspective. Athl Train Sports Health Care. 2016;8(4):145-153. 30. Honaker JA, Shepard NT. Use of the Dynamic Visual Acuity test as a screener for community-dwelling older adults who fall. J Vestib Res. 2011;21(5):267-276. 31. Ward BK, Mohammad MT, Whitney SL, Marchetti GF, Furman JM. The reliability, stability, and concurrent validity of a test of gaze stabilization. J Vestib Res. 2010;20(5):363-372. 32. Christy JB, Payne J, Azuero A, Formby C. Reliability and diagnostic accuracy of clinical tests of vestibular function for children. Pediatr Phys Ther. 2014;26(2):180-189. 33. Tsang WW, Wong VS, Fu SN, Hui-Chan CW. Tai Chi improves standing balance control under reduced or conflicting sensory conditions. Arch Phys Med Rehabil. 2004;85(1):129-137. 34. Hirsch MA, Toole T, Maitland CG, Rider RA. The effects of balance training and high-intensity resistance training on persons with idiopathic Parkinson’s disease. Arch Phys Med Rehabil. 2003;84(8):1109-1117. 35. Honaker JA, Converse CM, Shepard NT. Modified head shake computerized dynamic posturography. Am J Audiol. 2009;18(2):108-113. 36. Pang MY, Lam FM, Wong GH, Au IH, Chow DL. Balance performance in head-shake computerized

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48.

49.

dynamic posturography: aging effects and test-retest reliability. Phys Ther. 2011;91(2):246-253. Park MK, Lim HW, Cho JG, Choi CJ, Hwang SJ, Chae SW. A head shake sensory organization test to improve the sensitivity of the sensory organization test in the elderly. Otol Neurotol. 2012;33(1):67-71. Roma AA. Use of the head shake-sensory organization test as an outcome measure in the rehabilitation of an individual with head movement provoked symptoms of imbalance. J Geriatr Phys Ther. 2005;28(2):58-63. Clark S, Iltis PW. Effects of dynamic head tilts on sensory organization test performance: a comparison between college-age athletes and nonathletes. J Orthop Sports Phys Ther. 2008;38(5):262-268. Herdman SJ, Tusa RJ, Blatt P, Suzuki A, Venuto PJ, Roberts D. Computerized dynamic visual acuity test in the assessment of vestibular deﬁcits. Am J Otology. 1998;19(6):790-796. Peters BT, Mulavara AP, Cohen HS, SangiHaghpeykar H, Bloomberg JJ. Dynamic visual acuity testing for screening patients with vestibular impairments. J Vestib Res. 2012;22(2):145-151. Rine RM, Roberts D, Corbin BA, et al. New portable tool to screen vestibular and visual function: National Institutes of Health Toolbox initiative. J Rehabil Res Dev. 2012;49(2):209-220. Honrubia V, Bell TS, Harris MR, Baloh RW, Fisher LM. Quantitative evaluation of dizziness characteristics and impact on quality of life. Am J Otology. 1996;17(4):595-602. Perez N, Garmendia I, Garcia-Granero M, Martin E, Garcia-Tapia R. Factor analysis and correlation between Dizziness Handicap Inventory and Dizziness Characteristics and Impact on Quality of Life scales. Acta Otolaryngol Suppl. 2001;545:145-154. Gloor-Juzi T, Kurre A, Straumann D, de Bruin ED. Translation and validation of the vertigo symptom scale into German: A cultural adaption to a wider German-speaking population. BMC Ear Nose Throat Disord. 2012;12:7. Broglio SP, Ferrara MS, Sopiarz K, Kelly MS. Reliable change of the sensory organization test. Clin J Sport Med. 2008;18(2):148-154. Demer JL, Honrubia V, Baloh RW. Dynamic visual acuity: a test for oscillopsia and vestibulo-ocular reﬂex function. Am J Otology. 1994;15(3):340-347. McCrory P, Johnston K, Meeuwisse W, et al. Summary and agreement statement of the Second International Conference on Concussion in Sport, prague 2004. Phys Sportsmed. 2005;33(4):29-44. Dickin DC. Obtaining reliable performance measures on the sensory organization test: altered

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testing sequences in young adults. Clin J Sport Med. 2010;20(4):278-285. 50. Mishra A, Davis S, Speers R, Shepard NT. Head shake computerized dynamic posturography in peripheral vestibular lesions. Am J Audiol. 2009;18(1):53-59. 51. Vital D, Hegemann SC, Straumann D, et al. A new dynamic visual acuity test to assess peripheral vestibular function. Arch Otolaryngol--Head and Neck Surgery. 2010;136(7):686-691. 52. Schubert MC, Migliaccio AA, Della Santina CC. Dynamic visual acuity during passive head thrusts

in canal planes. J Assoc Res Otolaryngology. 2006;7(4):329-338. 53. Tian JR, Shubayev I, Baloh RW, Demer JL. Impairments in the initial horizontal vestibuloocular reďŹ&#x201A;ex of older humans. Exp Brain Res. 2001;137(3-4):309-322. 54. Tian JR, Shubayev I, Demer JL. Dynamic visual acuity during yaw rotation in normal and unilaterally vestibulopathic humans. Ann New York Acad Sci. 2001;942:501-504.

Appendix 1. Personal History Form Used in the Current Study

Please provide us with the important background information on the following form. If you do not understand a question, leave it blank and we will assist you. Thank you! NAME:________________________________________________________________ Sports you have participated in:____________________________________________ Occupation:____________________________________________________________ Reading level:__________________________________________________________ Have you EVER been diagnosed as having any of the following conditions? Circle the appropriate answer. YES NO Head injury If yes, how many prior head injuries?________ When was your last head injury? _____________________________________ YES NO Learning disorder YES NO Attention DeďŹ cit Disorder YES NO Dyslexia YES NO Heart Problems YES NO High blood pressure YES NO Circulation problems YES NO Asthma YES NO Bronchitis YES NO Thyroid problems YES NO Diabetes YES NO Depression YES NO Epilepsy or seizures YES NO Upper extremity musculoskeletal condition YES NO Lower extremity musculoskeletal condition YES NO Vertigo YES NO Visual problems Please describe if YES____________________________________

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Please list any surgeries or other conditions for which you have been hospitalized, including the approximate date and reason for the surgery or hospitalization: DATE Reason for Surgery/hospitalization 1._________________________________________ 2._________________________________________ 3._________________________________________ 4._________________________________________ 5._________________________________________ 6._________________________________________ Please describe any signiďŹ cant injuries for which you have been medically treated (including fractures, dislocations, sprains) and the approximate date of injury: DATE _______ _______ _______ _______

INJURY ___________________________________________________________ ___________________________________________________________ ___________________________________________________________ ___________________________________________________________

Please list any PRESCRIPTION medication you are currently taking __________________________________________________________________________________________ __________________________________________________________________________________________ __________________________________________________________________________________________ __________ How much caffeinated coffee or caffeine containing beverages do you drink per day?_______________ Do you smoke?_________How many packs of cigarettes do you smoke a day?____________________ Appendix 2. Symptom Inventory Scale Adapted from the UCLA DQ-2: Intensity of Dizziness Symptom Scale

0 No symptoms

1 Very mild

2 Mild

3 Moderate

4 5 Moderately Severe severe To assess symptoms of dizziness, each participant was asked after each test to rate their symptoms of dizziness from 0 to 5.

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IJSPT

ORIGINAL RESEARCH

COMPARISON OF A HEAD MOUNTED IMPACT MEASUREMENT DEVICE TO THE HYBRID III ANTHROPOMORPHIC TESTING DEVICE IN A CONTROLLED LABORATORY SETTING Eric Schussler, PhD, PT, ATC1 David Stark, MS2 John H. Bolte, IV, PhD2 Yun Seok Kang, PhD2 James A. Onate, PhD, ATC, FNATA2

ABSTRACT Background: Reports estimate that 1.6 to 3.8 million cases of concussion occur in sports and recreation each year in the United States. Despite continued efforts to reduce the occurrence of concussion, the rate of diagnosis continues to increase. The mechanisms of concussion are thought to involve linear and rotational head accelerations and velocities. One method of quantifying the kinematics experienced during sport participation is to place measurement devices into the athlete’s helmet or directly on the athlete’s head. Purpose: The purpose of this research to determine the accuracy of a head mounted device for measuring the head accelerations experienced by the wearer. This will be accomplished by identifying the error in Peak Linear Acceleration (PLA), Peak Rotational Acceleration (PRA) and Peak Rotational Velocity (PRV) of the device. Study Design: Laboratory study. Methods: A helmeted Hybrid III 50th percentile male headform was impacted via a pneumatic ram from the front, side, rear, front oblique and rear oblique at speeds from 1.5 to 5 m/s. The X2 Biosystems xPatch® (Seattle, WA) sensor was placed on the headform’s right side at the approximate location of the mastoid process. Measures of PLA, PRA, PRV from the xPatch ® and Hybrid III were analyzed for Root Mean Square Error (RMSE), and Absolute and Relative Error (AE, RE). Result: Seventy-six impacts were analyzed. All measures of correlation, fixed through the origin, were found to be strong: PLA R2=0.967 p<0.01, PRA R2=0.933 p<0.01, PRV R2=0.999 p<0.00. PLA RMSE was 34%, RE 31.0%±14.0, and AE 31.1%±13.7. PRA RMSE was 23.4%, RE -6.7±22.4 and AE 18.9%±13.8. PRV RMSE was 2.2%, RE 0.1±2.2, and AE 1.8±1.3. Conclusion: Without including corrections for effect of skin artifact, the xPatch® produces measurements highly correlated with the gold standard yet above the average error of testing devices in both PLA and PRA, but a low error in PRV. PLA measures from the xPatch® system demonstrated a high level of correlation with the PLA data from the Hybrid III mounted data collection system. Level of Evidence: 3 Key words: Concussion, head acceleration, head velocity

Old Dominion University, Norfolk, VA, USA The Ohio State University, Columbus, OH, USA

1 2

Disclosure None Acknowledgements We thank Sam Goldman, Alex Redrow, Arrianna Willis for testing setup and data acquisition.

CORRESPONDING AUTHOR Eric Schussler Old Dominion University 3107 Health Sciences Bldg, Norfolk, VA 23529 E–mail: ESchussl@odu.edu

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INTRODUCTION Reports estimate that between 1.6 and 3.8 million cases of concussion occur in sports and recreation each year in the United States.1 Despite continued efforts to reduce the occurrence of concussion, the rate of diagnosis continues to increase. The rate of occurrence has increased from 0.23 to 0.51 per 1000 high school athlete exposures in the period from 2005-2006 to 2011-2012.2 The cause of these concussions is often contact of a player’s head with the opponent’s head, body, or the ground.3 Sports related concussions affect more than 5% of high school and collegiate football players. Nearly 15% of the affected population goes on to sustain repeat concussions within the same season.4 Concussion has been identified as a potential risk factor for neurodegenerative dementia and decreased neurocognitive performance.5,6 Researchers utilize measurement devices to quantify the head impacts experienced by players in order to improve safety of athletes, but the technology of these devices often moves faster than the ability to independently test their accuracy. A method of quantifying head impulses experienced during sport is to place measurement devices such as accelerometers and gyroscopes into the athlete’s helmet or directly on the athlete’s head. Such technology allows potentially injurious accelerations to be quantified immediately during participation as well as for the collection of cumulative impact data from multiple players over the course of a season.7,8 Current wearable systems, often worn on or in helmets, have shown to be ineffective in accurately measuring the kinematics of the head.9–15 The X2 Biosystems’ xPatch® system (Seattle, WA) is an option to measure head kinematics directly from the scalp. This system consists of a triaxial linear accelerometer and a triaxial gyroscope contained in a 1½ in by ½ in device that attaches directly to the skin over the mastoid process of the athlete (Figure 1A). The system allows measurement of head accelerations for activities whose participants do not typically wear headgear or helmets. Previous research on the accuracy of a mouthguard based system manufactured and tested by the same company indicated the design may be capable of accurately measuring the accelerations experienced by the head16,17 yet limited research currently exists on the ability of the xPatch® device to accurately measure head accelerations.

The purpose of this study is to determine the measurement error of the xPatch® system when compared to a gold standard Hybrid III 50th percentile male headform (HIII) in a laboratory setting. Accurate measurements of the kinematics of the head during athletic competition are important in determining the risk of concussion, alerting medical personnel of the need for secondary evaluation and for developing concussion prevention strategies and equipment. It is theorized that these devices will have error percentages equivalent to other equipment currently commercially available. METHODS A helmeted HIII headform (Humanetics, Plymouth MA) was impacted via a pneumatic ram at varying impact speeds and directions. The HIII headform was attached to a HIII neck secured to a 40.23 kg mass on roller bearings which approximates the mass of a human thorax.18 The head was level (i.e. 0of tilt) for all impacts. The xPatch® sensor was placed on the headform’s right side at the approximate location of the mastoid process. Orientation of the sensor was set with the front of the device oriented perpendicular to the plane of testing (Figure 1A). The sensor was attached utilizing the manufacturer provided adhesive patch. The headform was then fitted with a Schutt Stallion lacrosse helmet (STX, Model: Stallion 500, Baltimore, MD) under which a wig comprised of human hair was placed and kept moist with a spray bottle, to simulate sweat. A Cascade STX chinstrap (Cascade Sports, Liverpool, NY) was used to secure the helmet to the headform and the helmet was fitted using manufacturer’s recommendations. A pneumatic ram weighing 23.9 kg was utilized to impact the helmet (Figure 1B). The impacting surface of the ram was an 8.25 inch diameter cylinder made of ultrahigh molecular weight polyethylene (UHMWPE). The helmet was impacted at four different speeds: 1.5, 2.5, 3.75, and 5 m/s from five different locations: frontal, side, rear, front-oblique and rear-oblique (Figure 2) to create a kinematic profile of impacts previously described to be representative of impacts experienced during play in light helmeted sports.14,19 Ram speed was controlled to within 0.1m/s. These speeds of impact were shown to produce accelerations at the head which have been reported during sports participation.10,20,21

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Figure 1. A) Front-Oblique testing setup. B) Affixation of the xPatch Device to the HIII.

Figure 2. Testing Matrix.

All oblique impacts were directed 45 from the midsagittal plane. Each impact was directed through the center of gravity of the headform. Four impacts were performed at each selected speed for each selected direction. The number, location and velocity of the impacts were controlled rather than resultant kinematics.

Instruments, Austin TX) where it was zeroed and filtered using CFC 1000 filters as per standard practices described in the Society of Automotive Engineers standard J211-1.23 While angular accelerations were not directly measured within the headform, they were calculated using algebraic equations as described by Padgaonkar.22

Linear accelerations and angular velocities of the HIII were recorded utilizing sensors placed within the headform. Piezoelectric accelerometers (Megitt’s Endevco, Model #:7246C-2000, Irvine CA) were used to measure linear acceleration while Angular Rate Sensors (ARS) (DTS ARS P18K Pro, Seal Beach CA) were used to measure angular velocity. Accelerometers were organized within the headform in a 3-2-2-2 configuration described by Padgaonkar22 typical for acquiring head impact data. Linear accelerations and angular velocities were stored using data acquisition system TDAS G5® (DTS, Seal Beach CA). Data were imported into DIAdem (National

Immediately prior to data collection, the xPatch® units were synced with the data recording system which set the time stamp on each device to the computer generated time. This same computer was utilized to run to HIII data acquisition system. Data from the xPatch® sensor were downloaded from the device into the manufacturer’s software at the completion of testing. This software automatically converts the linear acceleration data from the lateral aspect of the head ( a p ) to the center of gravity (aCG) (Equation 1) where  is angular velocity,  is angular acceleration and rp→CG is the geometrical relationship between point P, the location of the

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device on the head and Q, the location of the center of gravity of the head.24 aCG = a p +  × ( × rp→CG ) × rp→CG

(Eq. 1)

These calculations take place within the system utilizing manufacturer preset rp→CG distance and orientation which are not altered by the user. All data were transferred to excel spreadsheets to coordinate time stamps. The impacts measured by the xPatch® and sensors within the HIII were matched utilizing the time stamp on the xPatch® device and the data acquisition system regardless of the system determination of an actual impact or a non-intentional impact. The xPatch® system identifies false impacts through two methods including comparison to a set of wave form parameters and comparison to a reference waveform using cross-correlation.25 Two hundred and forty eight impacts were recording during testing. STATISTICAL METHODS Correlational analysis of peak linear acceleration, peak rotational acceleration and velocity were computed utilizing SPSS version 24 (IBM Armonk, NY). Error between the devices was calculated in Percent Relative Error (Equation 2), Percent Absolute Error (Equation 3) and Root mean square error (RMSE) (Equation 4) for each of the peak accelerations and velocities14. Percent Relative Error = HIII Measure − xPatch Measure *100 HIII Measure

(Eq. 2)

Percent Absolute Error = Hlll Measures − xPatch Meeasure *100 Hlll Measure RMSE =

∑

N i

( Relative Error )2 N

(Eq. 3)

(Eq. 4)

RESULTS Seventy-six tests were utilized for comparison of the xPatch® and HIII systems. Front and front oblique low speed tests included two impacts each that were below the 10g threshold of the xPatch® system as confirmed by the HIII system, thus they have been

excluded from analysis. All remaining tests were identified as impacts by the xPatch® system and were not determined to be non-impact signals referred to by the manufacturer as “clacks”. Resultant peak linear accelerations from the HIII headform ranged from 7.1 to 134.5g (average= 40.4g ±27.5), resultant peak rotational accelerations ranged from 606.8 to 8328.6 rad/s2 (average=2862.9 rad/s2 ±1889.2) and resultant peak rotational velocity ranged from 7.5 to 42.5 rad/s (average= 22.7 rad/s ±9.7). Analysis of the data indicated high correlations between the xPatch® and the HIII system. A correlation of linear acceleration fixed through the origin was found to be strong (p<0.01, R2=0.967). The results of rotational velocity were strongly correlated (p<0.00, R2=0.999) along with rotational acceleration (p<0.01, R2=0.933) when regressed through the origin (Figure 3). Analysis of the RE, AE and RMSE for Linear, and Rotational Acceleration and Rotational Velocity are presented in Table 1. RMSE for PLA in all combined directions was 34%, 2.8% for PRV and 23.4% for PRA. Percent relative error by averaged HIII and xPatch® measurements indicate an average error of 31% ±14.1 PLA, -6.7% ±22.6 PRA and 1.7% ±2.2 PRV (Table 1). Bland-Altman Plots are presented (Figure 4) for the percent error of each of the measures at the average measure of each of the devices. Average percent error for PLA was 31.0 (Limit of Agreement; 58.6, 3.3), average percent error for PRA was -6.7 (Limit of agreement: 37.6, -51.0), indicating an under estimation of the acceleration by the xPatch® system and average percent error for PRV was 1.7 (Limit of agreement: 6.0, -2.6). Significantly higher error was found between devices in RE and AE between the oblique measures over the non- oblique measures of PLA (Figure 5). Significant differences in AE: Front to Front Oblique (p=0.029), Front to Rear Oblique (p<0.001), Front-Oblique to Rear (p=0.018), Side to Rear Oblique (p=0.010), Rear Oblique to Rear (p<0.001). Significant Differences in RE: Front to Rear Oblique (p<0.001), Front-Oblique to Rear

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Table 1. Error by Type and Measurement

Figure 4. Bland-Altman Plot of Percent Error in A) Linear Acceleration B) Rotational Acceleration C) Rotational Velocity

Figure 3. Correlation of Measurements from xPatch to HIII system.

(p=0.044), Side to Rear Oblique (p=0.016), Rear Oblique to Rear (p<0.001). Significant differences were found in PRV between rear oblique measures and all other measures in RE and AE at p<0.001 (Figure 6). No significant difference was found

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between AE in PRA (p=0.199), with the only difference in RE between Rear Oblique and Rear (p=0.05, Figure 7). DISCUSSION As has been previously reported,15,26,27 there is a high correlation between the measurements of linear acceleration from a gold standard system and linear acceleration reported by the xPatch® system yet higher pooled RE and AE than other devices previously studied.15 All xPatch® measures of lin-

ear acceleration over-estimated the linear acceleration recorded by the HIII system. While pooled rotational acceleration measures were comparable to previously reported error seen in the X2 mouth guard and the Head Impact Telemetry® (HIT) System (Simbex, Lebanon, NH),15 the rotational velocity measures were quite accurate when compared to both the gold standard and other tested systems on market. Whereas linear accelerations were routinely over estimated, the RMSE of rotational acceleration from the xPatch® was found to be 24%, with errors

Figure 5. Percent Error in Linear Acceleration

Figure 6. Percent Error in Rotational Velocity.

Figure 7. Percent Error in Rotational Acceleration. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 597

Table 2. Percent Root Mean Square Error (RMSE) Measures by Direction and Magnitude of Impact

in both the positive and negative direction. The lack of uniformity in the error makes this measure difficult to interpret when measuring on the field accelerations, limiting clinical utility of the measure. The rotational velocity measure however was very accurate at 2.8% RMSE. The rotational velocity of the head is utilized for calculation of the Brain Injury Criterion (BrIC) which has strong support to be a predictor of brain injury in other applications such as motor vehicle accidents28. When analyzing the impacts from differing directions, oblique measures had higher error between the HIII and xPatch® measures for PLA and PRV. This response precludes the PLA and PRA data from being utilized to quantify peak acceleration of a single impact but does allow continued use of the pooled data to identify impact trends.

By directly placing the xPatch® on the head, this type of system may minimize the effect of interaction that has been found between the helmet and the head. Despite this, the design of the xPatch® has been shown to allow extraneous motion due to skin artifact. Accelerometers mounted on or within the helmet experience accelerations much larger than those that are experienced at the head.29 Athletes who do not wear tightly fitting equipment30 increase the disconnection between the accelerometers and the head, potentially introducing error.13 Studies identifying the motion artifact of the device during in vivo use found the device displaced on average 4mm from reference with the skull.26 This may affect the interpretation of these results, as the HIII skin may not accurately recreate the motion artifact of human skin. By placing the monitoring system directly on the head, the measure-

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ments are unaffected by the interaction of the impact and helmet design but the xPatch® unit has more extraneous motion than a mouthpiece based system.26 Identification of head impact severity utilizing head mounted and helmet mounted accelerometers and gyroscopes has been indicated to be unable to produce an accurate measurement of the forces experienced. The studies discussed throughout this paper indicate the usefulness of these devices may not lie in detecting a single impact injury but in quantifying the number of impacts that occur or utilization of pooled data. Future research should be performed to identify the effectiveness of these devices in monitoring total head impact exposure and the effectiveness of combined video review and impact profiles to reduce overall head contact. Limitations Limitations of this study include the absence of a mechanism to account for the difference in displacement of the device on human skin compared to the covering of the ATD. The xPatch® device is placed over the mastoid process and may displace up to 4mm26 introducing error to the measure that was not replicated in this study. The data collected contains more data points at lower accelerations. While this profile is representative of the effects of impacts expected in participation, additional testing should include higher acceleration tests. This research utilized one style of helmet throughout; different styles of helmets may change the acceleration profile which can be interpreted differently by the device. Because this device was tested specifically in conjunction with lacrosse helmets, the impactor design and speeds were best suited to recreate the acceleration profile involved in light helmeted sports.19,31 Additional work should be performed to identify the error source of this device through analysis of the raw data rather than the data as interpreted by the xPatch® software. This analysis may reveal a systematic error in the handling of the acceleration data by the software. Research should also identify the accuracy of derived measures of head acceleration such as Head Impact Criterion and Gadd Severity Index from the xPatch® compared to a gold standard measure. Future work should expand the range of impact speeds to be representative of the impacts experienced in different sports to accurately reflect the possible uses of the device.

CONCLUSIONS Accurate measurement of head accelerations experienced during sports participation is necessary for determining the specific mechanics of concussion in sport, determining methods to reduce concussions and identification of those who have experienced an impact that may have caused a concussion. The xPatch® System provides a strongly correlated overestimation of linear acceleration and a high level of accuracy in rotational velocity when compared to a gold standard measure. As linear acceleration is often the primary injury criteria used in sport at this time, consistent over estimation of the linear acceleration makes the xPatch® system a good tool to identify those in need of secondary injury screening by a qualified medical professional. REFERENCES

1. Langlois JA, Rutland-Brown W, Wald MM. The epidemiology and impact of traumatic brain injury: a brief overview. J Head Trauma Rehabil. 2006;21(5):375-378. 2. Rosenthal JA, Foraker RE, Collins CL, Comstock RD. National high school athlete concussion rates from 2005-2006 to 2011-2012. Am J Sports Med. 2014;42(7):1710-1715.3. 3. Zhang L, Yang KH, King AI. Biomechanics of neurotrauma. Neurol Res. 2001;23(2-3):144-156. 4. Guskiewicz KM, Weaver NL, Padua DA, Garrett WE. Epidemiology of concussion in collegiate and high school football players. Am J Sports Med. 2000;28(5):643-650. 5. Guskiewicz KM, Marshall SW, Bailes J, et al. Association between recurrent concussion and late-life cognitive impairment in retired professional football players: Neurosurgery. October 2005:719-726. 6. Stamm JM, Bourlas AP, Baugh CM, et al. Age of first exposure to football and later-life cognitive impairment in former NFL players. Neurology. 2015;84(11):1114-1120. 7. Broglio SP, Martini D, Kasper L, Eckner JT, Kutcher JS. Estimation of head impact exposure in high school football: Implications for regulating contact practices. Am J Sports Med. 2013;41(12):2877-2884. 8. Broglio SP, Eckner J, Martini D, Sosnoff JJ, Kutcher JS, Randolph C. Cumulative head impact burden in high school football. J Neurotrauma. 2011;28(10):20692078. 9. Cobb BR, Urban JE, Davenport EM, et al. Head impact Exposure in youth football: elementary

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school ages 9–12 years and the effect of practice structure. Ann Biomed Eng. 2013;41(12):2463-2473. 10. Crisco JJ, Wilcox BJ, Beckwith JG, et al. Head impact exposure in collegiate football players. J Biomech. 2011;44(15):2673-2678. 11. Daniel RW, Rowson S, Duma SM. Head impact exposure in youth football: Middle school ages 12 to 14 years. J Biomech Eng. 2014;136(9):094501. 12. Beckwith JG, Greenwald RM, Chu JJ. Measuring head kinematics in football: correlation between the Head Impact Telemetry System and Hybrid III headform. Ann Biomed Eng. 2012;40(1):237-248. 13. Jadischke R, Viano DC, Dau N, King AI, McCarthy J. On the accuracy of the Head Impact Telemetry (HIT) System used in football helmets. J Biomech. 2013;46(13):2310-2315. 14. Allison MA, Kang YS, Bolte JH, Maltese MR, Arbogast KB. Validation of a helmet-based system to measure head impact biomechanics in ice hockey: Med Sci Sports Exerc. 2014;46(1):115-123. 15. Siegmund GP, Guskiewicz KM, Marshall SW, DeMarco AL, Bonin SJ. Laboratory validation of two wearable sensor systems for measuring head impact severity in football players. Ann Biomed Eng. 2016;44(4):1257-1274. 16. Camarillo DB, Shull PB, Mattson J, Shultz R, Garza D. An instrumented mouthguard for measuring linear and angular head impact kinematics in American football. Ann Biomed Eng. 2013;41(9):19391949. 17. Hernandez F, Wu LC, Yip MC, et al. Six Degree-ofFreedom Measurements of Human Mild Traumatic Brain Injury. Ann Biomed Eng. December 2014. 18. Mertz HJ, Jarrett K, Moss S, Salloum M, Zhao Y. The Hybrid III 10-year-old dummy. Stapp Car Crash J. 2001;45:319-328. 19. Mihalik JP, Guskiewicz KM, Marshall SW, Blackburn JT, Cantu RC, Greenwald RM. Head impact biomechanics in youth hockey: comparisons across playing position, event types, and impact locations. Ann Biomed Eng. 2012;40(1):141-149. 20. Daniel RW, Rowson S, Duma SM. Head acceleration measurements in middle school football. Biomed Sci Instrum. 2014;50:291-296.

21. Broglio SP, Surma T, Ashton-Miller JA. High school and collegiate football athlete concussions: A biomechanical review. Ann Biomed Eng. 2012;40(1):37-46. 22. Padgaonkar AJ, Krieger KW, King AI. Measurement of angular acceleration of a rigid body using linear accelerometers. J Appl Mech. 1975;42(3):552. 23. Society of Automotive Engineers. J211-1 Instrumentation for impact test - Part 1 - Electronic instrumentation, in surface vehicle recommended practice. 1995. 24. X2Biosystems. Detection of head accelerations with an instrumented mouthguard and virtual projection sensors. 2011. 25. King D, Hume PA, Brughelli M, Gissane C. Instrumented mouthguard acceleration analyses for head impacts in amateur rugby union players over a season of matches. Am J Sports Med. 2015;43(3):614624. 26. Wu LC, Nangia V, Bui K, et al. In vivo evaluation of wearable head impact sensors. Ann Biomed Eng. 2016;44(4):1234-1245. 27. Nevins D, Smith L, Kensrud J. Laboratory evaluation of wireless head impact sensor. Procedia Eng. 2015;112:175-179. 28. Takhounts EG, Craig MJ, Moorhouse K, McFadden J, Hasija V. Development of brain injury criteria (BrIC). Stapp Car Crash J. 2013;57:243-266. 29. Manoogian S, McNeely D, Duma S, Brolinson G, Greenwald R. Head acceleration is less than 10 percent of helmet acceleration in football impacts. Biomed Sci Instrum. 2006;42:383-388. 30. Greenhill DA, Navo P, Zhao H, Torg J, Comstock RD, Boden BP. Inadequate helmet fit increases concussion severity in american high school football players. Sports Health Multidiscip Approach. 2016;8(3):238-243. 31. Lincoln AE, Caswell SV, Almquist JL, Dunn RE, Hinton RY. Video incident analysis of concussions in boys’ high school lacrosse. Am J Sports Med. 2013;41(4):756-761.

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IJSPT

ORIGINAL RESEARCH

TRANSVERSUS ABDOMINIS ELASTICITY DURING VARIOUS EXERCISES: A SHEAR WAVE ULTRASOUND ELASTOGRAPHY STUDY Kuniaki Hirayama1 Ryota Akagi2 Yuki Moniwa3 Junichi Okada1 Hideyuki Takahashi4

ABSTRACT Background: Although the transversus abdominis (TrA) is considered to play a significant role in maintaining trunk stability, there is little information regarding the type of exercise that best facilitates the development of tension in the TrA. Muscle elasticity shows a strong association with muscle tension. Shear wave ultrasound elastography provides a means by which the tension of TrA can be noninvasively estimated, by quantifying itâ&#x20AC;&#x2122;s elasticity. Purpose: The purpose of this study was to examine the TrA elasticity during several exercises as measured by shear wave ultrasound elastography, and to determine which of the studied exercises demonstrated the greatest tension. Methods: Ten healthy men performed abdominal hollowing, abdominal bracing, a hanging deadlift, elbowâ&#x20AC;&#x201C;toe plank with contralateral arm and leg lift, and back bridge with single leg lift. During these exercises, TrA elasticity was measured using ultrasound elastography. The same measurements were performed at rest before and after these exercises. Result: No significant difference was found for rest conditions measured before and after the exercises (p = 0.63). Abdominal bracing showed a significantly higher elasticity value than the other exercises (p < 0.05), except for hanging deadlift. Conclusion: Among the exercises, abdominal bracing was the exercise that elevated the TrA tension the most. The present results also suggested that hanging deadlift also produced comparably high TrA tension with abdominal bracing. Level of Evidence: 2c Key words: abdominal hollowing, abdominal bracing, deadlift, bridge exercise

Faculty of Sport Sciences, Waseda University, Saitama, Japan College of System Engineering and Science, Shibaura Institute of Technology, Saitama, Japan 3 School of Sport Sciences, Waseda University, Saitama, Japan 4 Department of Sports Science, Japan Institute of Sports Sciences, Tokyo, Japan 2

CORRESPONDING AUTHOR Kuniaki Hirayama, PhD Faculty of Sport Sciences Waseda University Mikajima 2-579-15, Tokorozawa, Saitama 359-1192, Japan Phone: +81-4-2947-6908 Fax: +81-4-2947-6908 E-mail: hirayama@aoni.waseda.jp

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INTRODUCTION Abdominal wall muscles function synergistically to maintain trunk stability. Because contraction of the transversus abdominis muscle (TrA) increases the tension of thoracolumbar fascia1 and intraabdominal pressure,2 TrA is considered to play a significant role in trunk stabilization.3 The abdominal hollowing or drawing in the abdominal wall is commonly used to facilitate TrA activation, particularly during the early-stage rehabilitation in patients with low back pain. In contrast, abdominal hollowing establish less trunk stability than that by abdominal bracing or cocontraction of abdominal wall muscles.4 The results of two electromyographic studies5,6 demonstrated that the TrA contracts during abdominal bracing. However, it is possible that other exercises, such as lifting and bridging exercises also activate the TrA to provide trunk stability. Studies that examine the effectiveness of a variety of exercises to activate the TrA are few.7 Such data would provide basic information in clinical rehabilitation or athletic training in order to promote better understanding of the role of the TrA in trunk stabilization. Because TrA is located in the deepest layer of the abdominal wall, it is difficult to estimate its activity or tension using surface electromyography or a myometer. Fine wire electromyography causes discomfort in subjects, and thus, cannot be easily used. In B-mode ultrasonography, a change in muscle thickness is used as an index of muscle activation. However, changes in TrA thickness could be influenced not only by TrA contraction but also by adjacent internal abdominal oblique muscle contraction, which shares the thoracolumbar fascia and rectus sheath. Because of these hindrances associated with the available techniques, a novel method to quantify TrA activation or tension is required. Therefore, the authors measured muscle elasticity during exercise using shear wave ultrasound elastography, which shows a strong association with muscle tension.8,9 This modality can non-invasively capture the propagation velocity of mechanical vibrations (i.e., shear waves) with imaging analyses, in which faster velocity indicates greater elasticity. Moreover, shear wave elastography can capture the elasticity of deep layers of muscles.10 The purpose of this study was to examine the TrA elasticity during several exercises as measured by shear wave ultrasound elastogra-

phy, and to determine which of the studied exercises demonstrated the greatest tension. METHODS Ten men (mean age: 22 ± 1 years, mean height: 176.2 ± 7.0 cm, mean weight: 70.6 ± 12.1 kg) participated in this study. No subject had low back pain or any other injury when the experiment was conducted. Written informed consent was obtained from each subject. The study was approved by the local ethics committee. The subjects performed seven exercises (Figure 1) in random order. Each exercise was repeated six times. The time intervals between the trials were set in such a way that subjects did not experience subjective fatigue (approximately 1 min). Before and after these exercises, measurements of TrA elasticity were performed in a resting condition to confirm that muscle elasticity did not change throughout the experiment. For each condition, including the resting condition, the subjects were asked to naturally inhale, then hold their breath at the end of exhalation, and not move during the measurement (i.e. for approximately five seconds). Details of each exercise condition were follows: Rest: The subjects lay supine with their legs extended. Abdominal hollowing: The subjects lay supine and were instructed to draw in the navel maximally towards the spine without any motion of the pelvis and thorax. Abdominal bracing: The subjects lay supine and were instructed to activate the abdominal muscles maximally without hollowing the lower abdomen. Hanging deadlift: The subjects held a 60-kg barbell with the hip and knee in a flexed position. The barbell height was standardized at above the patella. No instruction about the activity of abdominal wall muscles was provided. Elbow–toe plank with contralateral arm and leg lift: The subjects performed a front plank on their elbow and toe and then horizontally lifted their contralateral arm and leg (e.g., left arm and right leg). No instruction about the activity of abdominal wall muscles was provided.

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Figure 1. Pictures of various exercises performed in this experiment. 1: Abdominal hollowing; 2: Abdominal bracing; 3: Hanging deadlift; 4: Elbow-toe plank with left arm and right leg lift; 5: Elbow-toe plank with right arm and left leg lift; 6: Back bridge with right leg lift; 7: Back bridge with left leg lift.

Back bridge with single leg lift: The subjects lay supine and lifted their pelvis with a single leg. The other leg was extended during the bridge. No instruction about the activity of abdominal wall muscles was provided. Muscle elasticity of the TrA at the target region was measured using shear wave ultrasound elastography images (Figure 2) obtained by an ultrasonic apparatus (Aixplorer, SuperSonic Imagine, Aix-enProvence, France). Before the exercises, the target region, which was the right side of the umbilicus and 2 cm inside the mid-axillary line, was marked with an indelible pen. An 11-Hz electronic convex probe (SuperCurved 6-1, SuperSonic Imagine, Aixen-Provence, France) was used and placed on the target region transverse to the long axis of the body (i.e., parallel to the line of the muscle fibers of the TrA).11 Immediately after (approximately in 2-3 s) attaching the ultrasound probe on the skin, the color distribution is unstable (i.e., partial color missing and/or change in color) in some cases. A single image that captured a stable color distribution at a

specific time was selected from the images stored in the ultrasonic apparatus to estimate muscle elasticity.11-13 In the image, a 3-mm circle was set near the center of the TrA; the muscle elasticity within the circle was automatically calculated. The apparatus measures shear wave speed and calculates the elasticity by the following equation: E = 3ρc2/1000 Where E: Young’s modulus (kPa), ρ: Tissue density (kg/m3), c: Propagation speed (m/s). All the measurements were performed by a trained operator. For each subject, it took approximately 50 min for the entire procedure of the experiment. As mentioned above, the trial was repeated six times for each condition. Of the six measured values, a group of four measurements showing the lowest coefficient of variation among the twelve possible groups was adopted, and the mean values were used for further analysis.11 The authors failed to obtain the muscle elasticity during abdominal bracing and hanging deadlift for one subject; data from the

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The intraclass correlation coefficients type 1, 4 [ICCs (1, 4)] were calculated to examine the reliability of muscle elasticity measurements. The relative reliability among four measurements was calculated using one-way analysis of variance (ANOVA) model with consistency. The strength of the correlation was interpreted as follows: an ICC of ≤0.20 was “slight,” 0.21–0.40 was “fair,” 0.41–0.60 was “moderate,” 0.61–0.80 was “substantial,” and ≥0.81 was “almost perfect.”14 One-way ANOVA was performed to compare the muscle elasticity among exercises, and if appropriate, Tukey-HSD test was used for posthoc analysis. In this comparison, the average value of two time points was used for the rest condition. The significance level was set at p < 0.05.

Figure 2. Typical example of shear wave ultrasound elastography images during rest and abdominal bracing conditions. EO: external abdominal oblique, IO: internal abdominal oblique, TrA: transversus abdominis.

remaining nine subjects were presented as averages for these exercises. Statistical analysis Descriptive data were presented as mean ± standard deviation. A Student’s paired t-test was used to test the differences in muscle elasticity during the resting conditions before and after the exercises.

RESULTS The descriptive data for TrA elasticity and ICCs are presented in Table 1. The ICCs of muscle elasticity for each condition were significant (p < 0.05). Except for the elbow–toe plank with right arm and left leg lift, all conditions had ICC values higher than 0.89, which was evaluated as “almost perfect”.14 In addition, there was no difference in TrA elasticity during the resting condition before (14.1 ± 6.5 kPa) and after the exercises (13.6 ± 6.3 kPa) (p = 0.63). TrA elasticity was significantly higher during abdominal bracing than during other exercises (p < 0.05), except during hanging deadlift (Figure 3). DISCUSSION In the present study, most ICCs of the four trials reached the “almost perfect” level. Even one condi-

Table 1. Descriptive data of the TrA elasticity, intraclass correlation coefﬁcients (ICCs) and 95% conﬁdence intervals (CIs) of each condition

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Figure 3. Transversus muscle elasticity during various conditions. Abdominal bracing showed signiﬁcantly higher value than the other exercises, except for hanging deadlift. * Statistically signiﬁcant difference with abdominal bracing.

tion that did not reach the highest evaluation (i.e., “almost perfect”) can be evaluated as “substantial”.14 Therefore, it can be assumed that the measurements used in the present experiment were conducted with high reliability. In addition, no change in tension was observed during the resting conditions before and after exercise performance. This result suggests that the muscle elasticity measured during exercises did not reflect the changes in mechanical properties because of repeated exercises but can be assumed as an appropriate index of muscle tension. TrA elasticity was significantly higher during abdominal bracing than during other exercises, except during the hanging deadlift. Grenier & McGill4 reported that higher trunk stability was observed during abdominal bracing compared with that during abdominal hollowing. For abdominal bracing, not only the TrA but also “global muscles,” such as the external abdominal obliques, are recruited and have been assumed to contribute to the associated higher trunk stability. Higher TrA tension occurs during abdominal bracing than during abdominal hollowing which could be related to greater trunk stability. The deadlift is a typical strength and conditioning exercise that forces trunk flexion. High TrA tension was observed during the hanging deadlift, which would increase the tension of thoracolumbar fascia1 and intraabdominal pressure,2 and could be a strategy to resist trunk flexion load. The authors did not provide any instructions about recruitment of abdominal wall muscles during the hanging deadlift.

Therefore, it is possible that healthy men can automatically increase the TrA tension during resisted lifting exercise without special intention being paid to consciously increasing TrA tension. The results of this study do not diminish the significance or importance of abdominal hollowing. Rather, the findings suggest that abdominal bracing or lifting exercises increased TrA tension to a greater extent than during abdominal hollowing. This difference may be significant during late-stage rehabilitation or return to sport training to enhance the capability to maintain trunk stability during higher demand tasks. Hodges and Richardson15 revealed that the onset of TrA activity was delayed in patients with low back pain, and this delay can be remediated with an abdominal hollowing intervention.16 Particularly for early-stage rehabilitation, it may be necessary to revise the timing/onset of TrA activity by utilizing abdominal hollowing. This study has several limitations. The authors did not set criteria other than low back pain or other injuries for the subjects’ recruitment (e.g., activity level). This may have resulted in slightly large standard deviations of muscle elasticity during bracing or hanging deadlift and may have made the difference between hanging deadlift and other exercises insignificant. The ultrasound probe was manually operated. Therefore, the authors cannot completely eliminate errors associated with the manual operation. However, because the high reliabilities of the measurements were confirmed, these slight errors would be negligible.

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CONCLUSIONS Shear wave ultrasound elastography measures of the transversus abdominis at rest and during exercise trials demonstrated substantial to almost perfect reliability. Among the exercises adopted in the present study, abdominal bracing provided the highest tension of TrA. The hanging deadlift also produced comparably high TrA tension with abdominal bracing.

REFEERENCES 1. Tesh KM, Dunn JS, Evans JH. The abdominal muscles and vertebral stability. Spine (Phila Pa 1976). 1987;12(5):501-508. 2. Cresswell AG, Oddsson L, Thorstensson A. The influence of sudden perturbations on trunk muscle activity and intra-abdominal pressure while standing. Exp Brain Res. 1994;98(2):336-341. 3. Hodges PW, Richardson CA. Feedforward contraction of transversus abdominis is not influenced by the direction of arm movement. Exp Brain Res. 1997;114(2):362-370. 4. Grenier SG, McGill SM. Quantification of Lumbar Stability by Using 2 Different Abdominal Activation Strategies. Archives of Physical Medicine and Rehabilitation. 2007;88(1):54-62. 5. Barnett F, Gilleard W. The use of lumbar spinal stabilization techniques during the performance of abdominal strengthening exercise variations. J Sports Med Phys Fitness. 2005;45(1):38. 6. Urquhart DM, Hodges PW, Allen TJ, Story IH. Abdominal muscle recruitment during a range of voluntary exercises. Man Ther. 2005;10(2):144-153. 7. Okubo Y, Kaneoka K, Imai A, et al. Electromyographic analysis of transversus abdominis and lumbar multifidus using wire electrodes during lumbar stabilization exercises. J Orthop Sports Phys Ther. 2010;40(11):743-750.

8. Sasaki K, Toyama S, Ishii N. Length-force characteristics of in vivo human muscle reﬂected by supersonic shear imaging. J Appl Physiol (1985). 2014;117(2):153-162. 9. Yoshitake Y, Takai Y, Kanehisa H, Shinohara M. Muscle shear modulus measured with ultrasound shear-wave elastography across a wide range of contraction intensity. Muscle Nerve. 2014;50(1):103113. 10. Shinohara M, Sabra K, Gennisson JL, Fink M, Tanter M. Real-time visualization of muscle stiffness distribution with ultrasound shear wave imaging during muscle contraction. Muscle Nerve. 2010;42(3):438-441. 11. Hirayama K, Akagi R, Takahashi H. Reliability of ultrasound elastography for the quantiﬁcation of transversus abdominis elasticity. Acta radiologica open. 2015;4(9):2058460115603420. 12. Akagi R, Takahashi H. Acute effect of static stretching on hardness of the gastrocnemius muscle. Med Sci Sports Exerc. 2013;45(7):1348-1354. 13. Akagi R, Takahashi H. Effect of a 5-week static stretching program on hardness of the gastrocnemius muscle. Scand J Med Sci Sports. 2014;24(6):950-957. 14. Landis JR, Koch GG. The measurement of observer agreement for categorical data. Biometrics. 1977;33(1):159-174. 15. Hodges PW, Richardson CA. Delayed postural contraction of transversus abdominis in low back pain associated with movement of the lower limb. Journal of spinal disorders. 1998;11(1):46-56. 16. Tsao H, Hodges PW. Immediate changes in feedforward postural adjustments following voluntary motor training. Exp Brain Res. 2007;181(4):537-546.

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IJSPT

ORIGINAL RESEARCH

THE COMPARISON OF THE LUMBAR MULTIFIDUS MUSCLES FUNCTION BETWEEN GYMNASTIC ATHLETES WITH SWAY-BACK POSTURE AND NORMAL POSTURE Elnaz Mahdavie, PT, MS1 Asghar Rezasoltani, PT, Phd2* Leila Simorgh, PT, Phd3

ABSTRACT Background: The prevalence of sway back posture (SBP) is very high among elite gymnasts. This posture may be partly due to the improper function of lumbar multifidus muscles (LMM) as lumbar stabilizers muscles. Purpose: The aim of this study was to compare the thicknesses of LMM measured at rest and during the contraction elicited during an arm lift between elite gymnasts with SBP and normal posture. Study Design: Observational, descriptive, comparative Methods: The participants consist of twenty gymnasts between the ages of 17 and 30 who had trained in gymnastics for more than ten years. They were assigned to two groups: SBP (n=10) and control (n=10). Posture analysis with grid paper and plumb line was performed for all subjects. The thickness of LMM on dominant side of spinal column was measured by a real-time ultrasound at five lumbar levels. The thickness of the LMM was measured both at rest and during the contraction elicited during an arm lift. The variation between the LMM thickness between the muscle at rest and muscle at the peak of contraction was regarded as LMM muscle function. Result: The thickness of LMM was less in SBP group than the control group at all lumbar segments. The variation in LMM thickness between the state of rest and muscle contraction was significantly less in athletes with SBP than controls when compared at all levels of the lumbar spine (p < 0.05). Conclusion: The function of LMM may be disturbed in athletes with SBP as demonstrated by decreased thicknesses of LMM found in gymnasts with SBP. Additionally, the thickness of the LMM as a strong antigravity and stabilizing muscle group was decreased during arm raising in gymnasts with SBP. Level of Evidence: 3a Key words: Gymnastics, lumbar multifidus, sway-back posture, ultrasound imaging

Student Research Committee. School of Rehabilitation Sciences, Shahid Beheshti University of Medical Sciences, Tehran, Iran. 2 Physiotherapy Research Centre, School of Rehabilitation, Shahid Beheshti University of Medical Sciences, Tehran, Iran. 3 Department of Physiotherapy, School of Rehabilitation, Shahid Beheshti University of Medical Sciences, Tehran, Iran.

CORRESPONDING AUTHOR Asghar Rezasoltani Physiotherapy Research Centre, School of Rehabilitation, Shahid Beheshti University of Medical Sciences, Damavand Avenue, across from Bo-Ali hospital, Tehran, Iran. E-mail: a_rezasoltani@sbmu.ac.ir

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INTRODUCTION Posture, which is the relative disposition of the body at any one moment, is composite of the different positions of the joints at that time.1 A “well balanced posture” is the result of an interaction between the musculoskeletal system, nervous system and contextual effects.2 On the other hand, the faulty posture is result of an imperfect relationship among various skeletal structures of body. In such an example, the body is balanced less efficiently over its base of support. Therefore, any limitation, imbalance, or misalignment of musculoskeletal structures will have a significant effect on the efficiency of movements.

A significant correlation between SBP and low back pain has been reported by O’Sullivan et al.7 The authors revealed that lumbar multifidus activity decreased while subjects were in sway standing. The lumbar multifidus muscles (LMM) are important muscles that function to stabilize the lumbar vertebrae and pelvis. Researchers have suggested that the size and activity of LMM decreased in people with low back pain.8,9 Smaller size of LMM and loss of stability has been observed in the lumbar spine of patients with LBP.5,10,11

Sway-Back Posture (SBP) is the most common deviation or faulty posture of sagittal alignment1,3 characterized by anterior translation of the pelvis and hip joints, beyond the center of gravity line, a flatted curve in the lumbar region and overextended hip and knee joints.2 In one study, it has been reported that 35% of young girls may be affected by sway back posture.4 In a study that specifically included gymnasts, 80% of female gymnasts demonstrated SBP, which suggests that this postural fault is highly prevalent in this population.5

The multifidi are the deepest spinal extensors that serve to produce tension and contract to stabilize against external loads. According to different authors, it has been shown that actions like arm lift or straight leg raising, increase the load on the spinal column and result in the contraction of the spinal stabilizer muscles including the lumbar multifidus.5,28 Moseley et al30 indicated that activation of different elements of LMM occur concurrently with voluntary arm lift. Therefore, arm lift was chosen as a functional and dynamic task for use in this study to examine the contraction of the lumbar multifidi.

Gymnastics is a competitive sport that involves a series of maneuvers requiring strength, flexibility, balance and high levels of motor control. It consists of many different styles including artistic, rhythmic, trampoline, tumbling, and aerobic gymnastics. Artistic gymnastics is the best-known style, which includes different events for men and women. Women’s events include the vault, the uneven bars, the balance beam and the floor exercises. In gymnastics, there are six basic maneuvers including the hollow, arch, tuck, straddle, pike, and lunge, which are the basis for many skills used in by artistic gymnasts. In gymnastics, SBP may be induced as a result of competition or training programs.6 In some case, this posture may improve the gymnasts’ ability at different levels of the competition. However, this posture may also predispose the gymnasts to an imbalance in spinal musculature, which when carried into performance may induce pain and disability. Low back pain is the most common complaint reported by 50% of young artistic gymnasts (11-19 years old), occurring as a result of curve changes in lumbar region.6

Ultrasound imaging (USI) is a painless, noninvasive, and real-time technique that can be used to examine muscle features like size, thickness, and cross section area (CSA). This technique can be utilized either at rest or during dynamic movements and it has been reported Real-time ultrasonography is a relatively simple and inexpensive method of accurately measuring muscle thickness.14 Hodges et al24 noted that LMM thickness changes in response to other muscles contracting. This change in muscle thickness between rest and contracted conditions is one measure of contraction ability. It has been shown that USI is able to image both superficial and deep muscles such as the abdominal, lumbar multifidus, and pelvis floor muscles with high reliability and validity.12,13,14,15 The USI method for imaging the lumbar multifidus muscles (LMM) can be conducted in both prone and side-lying positions and without any adverse effects on the results.16 Further, it has been confirmed that USI is a reliable and noninvasive method for recording activity and thickness changes of lumbar multifidus (MF) in low load range of dynamic functions.17

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The aim of this study was to compare the thicknesses of LMM measured at rest and during the contraction elicited during an arm lift between elite gymnasts with SBP and normal posture. It was hypothesized that thickness changes (between relaxed and active conditions) of lumbar multifidus muscles may be different between gymnasts with sway-back posture and gymnasts without sway-back posture.

col was explained to all participants and it was their option to participate in this study. After acceptance, they were given a questionnaire to fill out. Height and weight was measured and the values were used to obtain their BMI. Consequently, the two groups were matched by means of age, time training and anthropometric variables such as weight, height and body mass index (BMI= weight / squared height) (Table 1).

MATERIALS AND METHODS

Posture analysis In order to evaluate posture, a plumb line and grid paper was used according to the methods described by Kendall.2 A similar technique was used by Mulhearn5 and Dolphens19 in their studies examining posture. The distance between the grid paper on the wall and plumb line that hung from the roof was 45 centimeters. The researchers adjusted the plumb line to be parallel to with one of lines on the grid paper. Markers were placed on the seventh cervical spinous process, the middle of greater trochanter and the middle of lateral malleolus. Then the gymnasts were asked to stand between the grid paper and plumb line so that the plumb line passed the middle of the lateral malleolus marker. Their feet were in the natural toe out position with 15cm distance between 1st metatarsals for all subjects. Then they were asked to put equal weight on both feet. They held this position for three minutes (this is necessary to allow enough time for them to achieve their habitual posture). To standardize the head posture, the ear meatus was aligned with

Sample Twenty female gymnasts between the ages of 17-30 were recruited. They all competed in artistic gymnastics at the Iranian national championship level and were recruited from Iranian Gymnastics Federation. They were assigned to two groups: gymnasts with SBP (n=10) and gymnasts with normal posture (n=10). Age, duration of activity at an elite level, time spent training per week, and history of surgery/pain in lower back and pelvis were self-reported by all athletes. To be included in the SBP group the subjects could not be performing exercise therapy or training for daily activities (such as correcting standing posture or doing exercises in order to correct the SBP). The exclusion criteria were: 1) any history of surgery of the spine, lower extremity or pelvis, and 2) musculoskeletal disorders like spasm, tendonitis, sprain and strain, fracture or dislocation of the spine or pelvis. Prior to involvement in the study, the study proto-

Table 1. Characteristic of the sway-back posture and control groups expressed as mean and standard deviation (SD) and range (signiďŹ cance level p < 0.05)

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the marker and the gymnasts were asked to look at a marker that was placed on the wall. The researchers then observed their final posture and checked the plumb line. If C7 and the greater trochanter were behind the plumb line, the gymnast was placed in the sway-back posture group, all others were placed in the normal posture group.19 Lumbar lordosis measuring protocol In the current study, a flexible ruler (Intra-tester reliability of 0.92, Inter-tester reliability of 0.82 and validity of 0.91 when compared to X-ray)20 was used to measure the angle of lumbar lordosis. At first, the lumbar vertebrae were identified using iliac crest as a landmark. The angle was measured after three minutes to reach to their typical posture. The flexible ruler was placed on the middle of the lumbar region, conforming to the spinal posture of the athlete (Figure 1). The spinous processes of L1 and S2 vertebrae were marked on the ruler. Then without altering the

shape of the flexible ruler, the convex side of the ruler was drawn on a paper and L1 and S2 were delineated on the paper. To evaluate the lumbar lordosis two points of L1 and S2 were connected to each other and identified as “L” then from the middle of the line “L”, another line perpendicular to the curve was added and labeled it as “H”. The following formula was then used to compute the degree of lumbar lordosis (D).

D = 4[arc tan (2H/L)] Pelvic tilt measurement Pelvic tilt was measured with inclinometer. Arms of inclinometer were placed on the ASIS and PSIS landmarks. Then the inclination of the pelvic directly was read from protractor (Figure 2). To evaluate posterior or anterior pelvic tilt, vertical lines from anterior superior iliac spine (ASIS) and posterior superior iliac spine (PSIS) were drawn to grid paper. These two points were attached to each other with a line. The distance between the middle of line and plumb line was measured with a ruler.20 Intraclass correlation coefficient for repeated measures, Intra-tester and Intra-tester reliability were indicated 0.99.29 Ultrasound imaging protocol The technique was performed using a B-Mode ultrasonography device (Honda 2100, Honda Co., Japan)

Figure 1. Lumbar lordosis measurement using ﬂexible ruler. The spinous processes of L1 and S2 was marked on the ruler. (L1: ﬁrst lumbar vertebrae, S2: second sacrum vertebrae).

Figure 2. Pelvic tilt was measured with inclinometer. Arms of inclinometer were placed on the ASIS and PSIS landmarks. (ASIS: anterior superior iliac spine, PSIS: posterior superior iliac spine).

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with a frequency of 7.5 M Hz and a linear array probe. The athletes were placed in prone position with a pillow under their abdominal region to flatten their lordosis. They were asked to relax in order to prevent any muscle contraction. The researchers used the iliac crest as a landmark to find lumbar vertebrae, and the top of the crest at the spine was identified as L4. Palpation was then used to identify other lumbar vertebrae, which were delineated by marks on the skin. By putting the probe longitudinally, the spinous processes of the vertebrae were seen on screen. Once each vertebra was found, the probe was turned to a transverse alignment. In this view, the spinous process and both transverse processes of one vertebra were identifiable. The echogenic transverse process of the dominant side of the spine was used as a landmark. Dominance was defined by asking the subject which hand she would use to write. The distance between the transverse process and the subcutaneous fascia was used as lumbar multifidus muscle thickness while the subjects were at rest. Athletes were asked to elevate their dominant arm up to the level of their ear in order to induce a contraction of the LMM. At the end of the movement, the muscle thickness was measured (Figure 3). These subsequent steps and measurements were

repeated at each vertebral level from L1-L5. A total of 10 measurements were taken per gymnast to view each vertebral segment from L1-L5 while at rest and during contraction of the LMM. The thicknesses of lumbar multifidus muscles were measured on the display of the ultrasound device, then the thickness variation between the state of rest and contraction was computed as a measure of muscle contractility. Reliability The reliability of the LMM thickness, the degree of lumbar lordosis and the pelvis tilt measurement was evaluated for ten randomly selected individuals (five from SBP group and five from control group). The measurements were repeated three times with an interval of one hour within one day. The first measurement was performed as mentioned, and for subsequent measurements all markers were removed after each testing session. For all subsequent measures, bony landmarks were re-identified and markers were replaced as mentioned in the methods section. The level of ICC (3,1) for the intratesters reliability was excellent for LMM thickness measurement in both states of rest [ICC = 0.98 (95% CI = 0.92-0.99)] and contraction [ICC = 0.96 (95%

Figure 3. (a) Ultrasound image of lumbar multiďŹ dus muscle at the ďŹ fth level of vertebra with measurement of the muscle thickness at rest. (b) Ultrasound image of the lumbar multiďŹ dus contraction consequent to arm lift. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 611

CI = 0.88-0.98)]. Excellent values were also found for the flexible ruler [ICC = 0.99 (95% CI = 0.97-.099)]; and inclinometer [ICC = 0.97 (95% CI = 0.88-0.99)]. Statistical analysis: The Shapiro-Wilk test was used to assess for normal distribution and the Levene test was used to assess the equality of variances. An ANCOVA was used to assess outcomes using the covariate variables such as age, weight, height and BMI. Data analysis was then rendered using SPSS (version 23). Statistical significance was established as p< 0.05. RESULTS The measurement of lumbar lordosis was significantly less in gymnasts with SBP than the control group (p ≤ 0.008) demonstrating decreased lordosis in this posture. However, the pelvic tilt difference between the two groups was not statically significant different. The thicknesses of the lumbar multifidus muscles were less in the SBP group compared to the control group in all lumbar segments from L1-L5 at rest, but these differences were not statistically significant. However, the thickness during contraction as induced by the action of arm lift was significantly different (p ≤ 0.031) at the level of fifth lumbar vertebra (Tables 2 and 3). The LMM thicknesses during contraction minus the LMM thicknesses at rest indicated that the thickness variation was significantly less in gymnasts with

SBP when compared to the control group at all five segments (L1-L4, p < 0.05 and L5, p value < 0.001) (Table 4). The measurement of lumbar lordosis (measured with a flexible ruler) was significantly less in gymnasts with SBP than the control group (p ≤ 0.008) demonstrating decreased lordosis in this posture. However, the pelvic tilt difference between the two groups was not statistically significantly different. DISCUSSION In this study, USI was used to evaluate the LMM in gymnastic athletes with and without SBP. The thickness of LMM was measured in states of both rest and contraction. The variations of thickness between the two states are considered a measurement of LMM function. Sway-back posture is an adaptation of spine that occurs over time with a lot of potential side effects. It is not considered an acute injury such as those that include sprains, fractures, dislocations, muscle strains and contusions.21 Many different variables contribute to development of SBP, including the method of coaching and type of exercises, as well as growth and development. Postural adaptations and gymnast preference appear to be important factors among gymnastic athletes that develop SBP. Training volume is another important factor and it has been shown that training more than 400 hours/year may decrease or increase the degree of lumbar lordosis,22 hovever, there was no clear difference between gymnasts volume of training between those with SBP and those without. In the current study, the athletes with SBP focused more on the vault events which requires the straight straddle and tuck

Table 2. Thickness of lumbar multiﬁdus muscle in all lumbar segments from L1-L5 in rest in the sway-back and control groups expressed as mean and standard deviation (SD) and range (signiﬁcance level p < 0.05)

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Table 3. Thickness of lumbar multiﬁdus muscle in all lumbar segments from L1-L5 in contraction consequent to arm lift in the sway-back and control groups expressed as mean and standard deviation (SD) and range (signiﬁcance level p < 0.05)

Table 4. Thickness changes of lumbar multiﬁdus muscle between rest and contraction in response of dominant arm lifting in dominant side of spine column in all lumbar segments from L1-L5 in the sway-back and control groups expressed as mean and standard deviation (SD) and range (signiﬁcance level p < 0.05)

positions. Furthermore, they also performed exercises like handstand flat back and handstand backward rolling in order to be successful on the vault. Therefore, while focusing on these, it is possible that differences in evolution of spinal column posture may have occurred as a result of spinal muscles imbalances, that may have happened between agonist and antagonist or synergistic spinal muscles. Group comparison Variation in muscle size as measured by USI has been regarded as a sign of muscle efficacy, tension and

contractility.23,24 Muscle size, has been reported to significantly correlate with electromyography activities of the same muscle.24 Muscle thickness is a factor to determine the level of contractility generated by an individual muscle’s contraction.23,24 In paraspinal muscle ultrasonography, the thickness of semispinalis capitis muscle significantly increased as the force of the neck extensor muscles were increased from 0% to 100% of maximum voluntary contraction.23 In the current study, the thickness variation between state of rest and contraction (consequent to dominant

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arm flexion) was used as a measure of LMM function. The authors’ main goal was to determine whether these thickness changes were different between gymnasts with SBP and gymnasts with normal posture. To find the activity or contractility of muscles by USI, difference of thickness between rest and contraction is important. Less difference of measured muscle thickness between rest and contraction indicates less activity of the LMM23,24 because the LMM are enabled to stabilize the spinal column in a correct alignment during dynamic movements. Ultrasound imaging of LMM at all five segments demonstrated significant differences in their thickness variations (obtained from the LMM thicknesses during contraction induced by the action of arm lift minus the LMM thicknesses at rest) between the SBP group and control group. Significantly greater thickness variation of LMM between two groups at the level of fifth lumbar vertebra indicates that muscles at this level were influenced by SBP. This finding was in accordance with that found by Pezoloto et al25 who documented greater fat infiltration in lumbar multifidus and other erector spine muscles at the levels of fourth and fifth lumbar vertebrae in subjects with SBP. This measurement was used to confirm greater atrophy of muscle due to the presence of more fat and may be it is because of anatomical and biomechanical properties that the greatest volume of LMM is at the L5/S1 level. Findings in the current study indicate that during contraction, the LMM thickness in SBP group was significantly less than the control group. The amount of lumbar lordosis was significantly less in SBP subjects than the control group, which could be due to the stabilizer role of LMM. The LMM are important for maintenance of the lumbar lordosis and activity of the LMM contributes to efficient alignment of spinal column and pelvis, and those athletes with SBP had decreased activity during arm raise of the LMM. However, anterior pelvic tilt was not significantly different between the SBP subjects and the controls. Anterior pelvic tilt has been described as an important abnormality in SBP, which causes anterior shift of trunk to the line of gravity as the athlete tends to extend the trunk, with a corresponding decrease in the involvement of LMM and paraspinal muscles.25,26,27 On the other hand, the abdominal muscles such as rectus abdominis

increase their work to attempt to impede the extension of lumbar vertebrae.28 These muscle imbalances have been reported in a variety of different studies.3,15,29 However, in those studies, the athletes were asked to actively recreate different postures such as sway standing or sway sitting, whereas in the current study the gymnasts were asked to demonstrate their normal, habitual posture. This difference may explain the lack of significant difference in anterior pelvic tilt found between groups in the current study LIMITATIONS & RECOMMENDATIONS Only athletes who were in artistic gymnastic competition were evaluated and hence more general biomechanical tests were not performed. The current results may not apply to all gymnasts. There are other limitations like sample size, lack of blinding and only the female gender being represented. The examiner was not blinded to the either groups. The authors recommend investigating the biomechanical properties of the type of exercises and training programs, which may influence the development of SBP among gymnasts which could be beneficial for education for coaches and to guide treatment for physical therapists. CONCLUSION In this study, ultrasound imaging was used to evaluate the LMM in gymnastic athletes with SBP. The thickness of LMM was measured in both states of rest and contraction. The variations of thickness between the two states are considered a measurement of LMM contractility. The contractility of LMM was decreased in athletes with SBP in comparison with athletes with normal posture. This may affect the normal lumbar curvature or stabilization during activity as a result of insufficient LMM contraction. The authors recommend this method to assess and evaluate the function of LMM and spinal column posture in individuals with SBP. REFERENCES 1. Magee DJ. Orthopedic physical assessment. 5th edition. Saunders, Philadelphia, PA; 2007. 2. Kendall, FP, McCreary, EK, Provance PG, et al. Muscles: Testing and Function. 5th edition. Lippincott Williams & Wilkins, Baltimore, MD; 2005. 3. Smith A, O’Sullivan PB, Straker L. Classiﬁcation of sagittal thoraco-lumbo-pelvic alignment of the adolescent spine in standing and its relationship to low back pain. Spine.2008;33:2101-2107.

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4. Simorgh L, Kheirkhah M, Khakhali-Zavieh M. Investigation of sway back posture prevalence and alignment of spine and lower limb joints in this deformity. Rehabilitation J. 2006;7(2):31-37. 5. Mulhearn S, George K. Abdominal muscle endurance and its association with posture and low back pain: An initial investigation in male and female elite gymnasts. Physiotherapy J. 1999; 85(4)210-216. 6. Purcell L, Micheli L. Low Back Pain in Young Athletes. Sports Health. 2009; 1(3): 212-222. 7. Mitchell T, O’Sullivan PB, Burnett AF, et al. Regional differences in lumbar spinal posture and the influence of low back pain. BMC Musculoskel Disord. 2008;18(9):152-163. 8. Hodges PW, Moseley GL. Pain and motor control of the lumbopelvic region:Effect and possible mechanisms. J Electromyog KineS. 2003;13(4):361-370. 9. Lee SW, Chan CK, Lam C, et al. Relationship between law back pain and lumbar multifidus size at different postures. Spine. 2006; 31(19): 2258-2262. 10. Dickx N, Cagnie B, Achten E, et al. Differentiation between deep and superficial fibers of the lumbar multifidus by magnetic resonance imaging. European Spine J. 2010; 19(1):122-8. 11. Pavlova V, Meakin JR, Cooper K, et al. The lumbar spine has an intrinsic shape specific to each individual that remain a characteristic throughout flexion and extension. Eur Spine J. 2014; 23(1): 26-32. 12. Lee JP, Tseng WY, Shau YW, et al. Measurement of segmental cervical multifidus contraction by ultrasonography in asymptomatic adults. Manual Therapy J. 2007; 12(3): 286-94. 13. Kristjansson E. Reliability of ultrasonography for the cervical multifidus muscle in asymptomatic and symptomatic subjects. Man Ther. 2004; 9(2): 83-8. 14. Dupount AC, Sauerberi EE, Fenton PV, et al. Real time sonography to estimate muscle thickness: comparison with MRI and CT. J Clin Ultrasound. 2001;29(4):230-38. 15. Lee JC, Cha JG, Kim Y, et al. Quantitative analysis of back muscle degeneration in the patients with the degenerative lumbar flat back using a digital image analysis: comparison with the normal control. Spine. 2008; 33(3): 318-325. 16. Coldron Y, Stokes M, Cook K. Lumbar multifidus muscle size dose not differ whether ultrasound imaging is performed in prone or side lying. Man Ther. 2003;8(3):161-165. 17. Kiesel KB, Uhl TL, Underwood FB. Measurement of lumbar multifidus muscle contraction with rehabilative ultrasound imaging. Man Ther. 2007; 2(2):161-166.

18. McDonnell, MK, Sahrmann, SA, Van Dillen A specific exercise program and modification of postural alignment for treatment of cervicogenic headache: a case report. J Ortthop Sports Phys Ther. 2005;35(1):3-15. 19. Dolphens M, Cagnie B, Vleeming A. A clinical postural model of sagittal alignment in young adolescents before age at peak height velocity. Eur Spine J. 2012; 21(11): 2188-2197. 20. R Rajabi, F Seidi, F Mohamadi. Which method is accurate when using the flexible ruler to measure the lumbar curvature angle deep point or midpoint of arch. World Appl Sci. 2008;4(6):849-852. 21. Fettorne FA, Ricciardelli E. Gymnastic injuries: the Virginia experience 1982-1983. Am J Sports Med. 1987;15(1):59-62. 22. Jose M. Muyor, Pedro A. Lopez-Minarro, Fernando Alacid. Spinal posture of thorasic and lumbar spine and pelvic tilt in highly trained cyclists. J Sports Sci Med. 2011;10(2):355-361. 23. Rezasoltani A, Ylinen J, Vihko V. Isometric cervical extension force and dimension of semispinalis capitis muscle. J Rehab Res and Dev. 2002;39(3):423428. 24. Hodges PW, Pengel LHM, Herbert RD, Gandevia SC. Measurement of muscle contraction with ultrasound imaging. Muscle & Nerve. 2003;27(6):682-692. 25. Pezolato A, de Vasconcelos E, Defino HLA, NogueiraBarbosa MH. Fat infiltration in the lumbar multifidus and erector spine muscles in subjects with swayback posture. Eur Spine J. 2012;21(11):2158–2164. 26. O’Sullivan PB, Grahamslaw M, Kendell M, et al The effects of different standing and sitting postures on trunk muscle activity in a pain-free population. Spine. 2002;27(11):1238–1244. 27. Sahrmann SA. Diagnosis and treatment of movement impairment syndromes. 1st edition. Mosby, St. Louis, MO; 2002. 28. Bogduk N. Clinical anatomy of the lumbar spine and sacrum, 4th edition. Elsevier, Philadelphia, PA; 2005. 29. Claus AP, Hides J, Mosley GL, et al. Different way to balance the spine: subtle changes in sagittal spinal curves affect regional muscle activity. Spine. 2009;34(6):208-14. 30. Moseley GL, Hodges PW, Gandevia SC. Deep and superficial fibers of the lumbar multifidus muscle are differentialy active during voluntary arm movements. Spine. 2002; 27(2): E29-3. 31. Chowdhury Salian Sh, Yardi S. Intertester and Intratester Reliability and Validity of Measures of Innominate Bone Inclination in Standing Using Pelvin. Int J Innovations Sci Engineer Tech. 2015; 7(4): 5068-77.

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IJSPT

ORIGINAL RESEARCH

CAN RUNNERS PERCEIVE CHANGES IN HEEL CUSHIONING AS THE SHOE AGES WITH INCREASED MILEAGE? Mark W. Cornwall, PT, PhD, FAPTA1 Thomas G. McPoil, PT, PhD, FAPTA2

ABSTRACT Background: For those runners who utilize footwear and have a rearfoot strike pattern, the durability of the midsole heel region has been shown to deteriorate as shoe mileage increases. Purpose: The purpose of this study was threefold: 1) to determine if the runner can self-report changes in heel cushioning properties of the midsole after an extended period of distance running, 2) to determine if force and plantar pressures measured in the heel region of the midsole using a capacitance sensor insole change after running 640 km, and 3) to determine if a durometer could be used clinically to objectively measure changes in the hardness of the material in the heel region of the midsole. Study Design: Cross-sectional Study Methods: Fifteen recreational runners voluntarily consented to participate and were provided with a new pair of running shoes. Each participantâ&#x20AC;&#x2122;s running style was observed and classified as having a rearfoot strike pattern. Inclusion criteria included running at least 24 km per week, experience running on a treadmill, no history of lower extremity congenital or traumatic deformity, or acute injury six months prior to the start of the study. The ability of each participant to self-perceive changes in shoe cushioning, comfort and fit was assessed using the Footwear Comfort Assessment Tool (FCAT). In-shoe plantar pressures and vertical forces were assessed using a capacitance sensor insole while runners ran over a 42-meter indoor runway. A Shore A durometer was used to measure the hardness of the midsole in the heel region. All measures were completed at baseline (zero km) and after running 160, 320, 480, and 640 km. In addition to descriptive statistics, a repeated measures analysis of variance was used to determine if the FCAT, pressures, forces, or midsole hardness changed because of increased running mileage. Result: While plantar pressures and vertical forces were significantly reduced in the midsole heel region, none of the runners self-reported a significant reduction in heel cushioning based on FCAT scores after running 640 km. The use of a durometer provided an objective measure of the changes in the heel region of the midsole that closely matched the reductions observed in pressure and force values. Conclusion: The results indicated that runners who have a rearfoot strike pattern will have a 16% to 33% reduction in the amount of cushioning in the heel region of the midsole after running 480 km. Although there were significant reductions in heel cushioning, the experienced recreational runners in this study were not able to self-perceive these changes after running 640 km. In addition, the use of a durometer provides a quick and accurate way to assess changes in the hardness of the heel region of the midsole as running mileage increases. Level of Evidence: 3, Controlled laboratory study Key words: Durometer, force, midsole, plantar pressure, running

Department of Physical Therapy and Athletic Training, Northern Arizona University, Flagstaff, AZ 2 School of Physical Therapy, Regis University, Denver, CO

CORRESPONDING AUTHOR Thomas G. McPoil, PT, PhD School of Physical Therapy Regis University 3333 Regis Blvd., G-4 Denver, CO, 80221 Phone: 303 964-5137 E-mail: tmcpoil@regis.edu

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INTRODUCTION For those runners who utilize footwear when training, a major concern is the durability of the shoe midsole which serves to provide cushioning when impacting the supporting surface. Since previous research suggests that runners using shoes will have a greater tendency to have a rearfoot strike, the durability of the heel region of the midsole is of particular concern.1,2 Runners can experience impact forces of 2.5 times body weight when the shoe collides with the ground3 and typically there are approximately 600 to 750 foot strikes per km while distance running.4 Thus, the durability of the running shoe midsole is a major consideration when deciding when to purchase new footwear. The two most common materials used in midsole construction are ethylene vinyl acetate (EVA) and polyurethane (PU). One of the first studies to assess the deterioration of running shoe midsole materials was conducted by Cook, et al.5 In their study, a mechanical impact tester was used to assess midsole durability in addition to two experienced runners. They tested 13 different running shoes that all had midsoles constructed from EVA. They reported that with mechanical impact testing the running shoes tested retained less than 60% of the initial shock absorption capacity between 400 to 800 km of wear. The degree of midsole degradation that resulted when the two experienced runners used the same running shoes was only a 20% to 30% reduction between 480 and 640 km. In an attempt to overcome the differences between mechanical and in-vivo testing of midsole durability, Hamill and Bates assessed six healthy male runners using the same shoe with a dual-density EVA midsole while running over a force platform at running intervals separated by 140 km.6 All runners in the study ran in excess of 48 km per week for three years prior to the start of the study. Although they reported a loss of 7.3% in the shock absorbing capability of the shoe after running 420 km, the magnitude of the loss was much less than reported by Cook, et al. Hamill and Bates also noted that the runners’ ability to “sense” the performance characteristics of the shoe may be misleading since reasonable functional changes in the midsole occurred during the initial 300 to 400 km of wear. This inability of the runner to “sense” degradation of midsole cushioning could be a factor in the development of running related injuries since

Kong, et al have shown that as running shoe cushioning capability decreases, runners modify their running pattern to maintain constant external loads and that the adaptation strategies used by runners were not affected by different cushioning technologies (i.e.; air, gel).7 In a more recent study, Schwanitz and Odenvald assessed durability of the heel region of the midsole using mechanical impact testing and reported a 20% reduction in shock absorption characteristics of the midsole after simulating approximately 600 km of running.8 While Wang, et al assessed midsole cushioning at 50 km increments using a mechanical impact tester, they had eight male amateur runners run in the test shoes for 500 km.9 They reported a significant decrease in cushioning but the reduction was only 5%. In the only study to date that has attempted to use an in-shoe pressure sensing insole to assess midsole durability, Verdejo and Mills reported that plantar pressures in the heel region increased on average by 100% after three healthy males ran 500 km.10 All three runners were rearfoot strikers and utilized shoes with an EVA midsole. The increase in plantar pressures was attributed to fatigue of the EVA foam causing the material to become harder. In interpreting the research to date, it would appear that mechanical testing over-estimates the degradation of the cushioning properties of the EVA midsole when compared to in-vivo testing of midsole durability. Authors of in-vivo studies to date have reported that the degradation of midsole cushioning can occur after running anywhere between 480 to 640 km. In responding to a runners’ inquiry as to when they should buy a new pair of running shoes to ensure adequate cushioning, it is important for the clinician to know 1) if the runner can self-perceive a degradation in midsole cushioning, especially in the heel region if they are a rearfoot striker, and 2) whether there is a simple test that can be done in the clinic to assess possible degradation of the midsole. Mundermann, et al developed a Footwear Comfort Assessment Tool (FCAT) to allow the runner to selfreport their satisfaction with shoe comfort, fit and cushioning.11 The FCAT consists of nine 100 millimeter visual analogue scales that assesses overall comfort, heel cushioning, forefoot cushioning, pronation-supination control, arch height, heel cup

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fit, shoe heel width, shoe forefoot width, and shoe length. The left end of the all nine scales or zero (0 millimeters) was labeled “not comfortable at all” and the right end of the scale (100 millimeters) was labeled “most comfortable condition imaginable.” The higher the score for all nine scales the better the comfort, fit and cushioning. Mundermann, et al reported that the FCAT had excellent levels of reliability for all scales if a control condition was used before each session the FCAT was utilized.11 In attempting to find a simple tool that could be used in the clinic to assess the hardness of the EVA foam, Barton, et al reported excellent levels of reliability between three raters using a hand-held durometer to assess the hardness of the midsole in the heel region at the point where the center of the heel contacts the midsole within the shoe.12 While Barton, et al provide the clinician with a test to consistently measure the hardness of the midsole, it is unknown if assessing the hardness of the midsole in the heel region with a durometer would be sensitive enough to assess degradation of the midsole over an extended period of running. After an extensive review of the current literature, the authors could not find any studies that have assessed runners’ self-perceived changes in heel cushioning provided by the midsole or whether a durometer to test midsole hardness in the heel region of the midsole would be effective over an extended period of distance running. Thus, the purpose of this study was threefold: 1) to determine if the runner can self-report changes in heel cushioning properties of the midsole after an extended period of distance running, 2) to determine if force and plantar pressures measured in the heel region of the midsole using a capacitance sensor insole change after running 640 km, and 3) to determine if a durometer could be used clinically to objectively measure changes in the hardness of the material in the heel region of the midsole. Three hypothesizes were developed for this study. First, those individuals who run at least 24 km per week would be able to perceive a reduction in the heel cushioning when running in the same pair of shoes over a running distance of 640 km (approximately 400 miles). Second, that the use of a durometer could be used to objectively measure changes in midsole material hardness in the heel region of running shoes used

by the same runner over a running distance of 640 k. Third, that reductions in force and plantar pressures in the rearfoot or heel region of the midsole would not be greater than between 20% and 30% after running 640 km. METHODS Subjects Fifteen recreational runners (4 male; 11 female) with a mean age of 26.3 (sd=4.4) years volunteered to participated in this study. All participants were recruited from the greater Flagstaff, Arizona, region using advertisements in printed media and notice boards. All runners selected had no previous history of surgery, childhood or congenital disorders, fractures or dislocations to the lumbar spine, lower extremity, ankle or foot. In addition, none of the runners had a history of a trauma or pain to either the lower extremity, ankle and foot, or lumbosacral regions for at least six months prior to start of the investigation. All the participants had consistently run at least 24 km per week for two years prior to the start of the study. Each participant’s running style was visually observed by one of the investigators (TGM) and all 15 runners were classified as having a rearfoot strike pattern. The Northern Arizona University Institutional Review Board approved the study procedures and each participant signed an informed consent prior to taking part in the study. After signing the informed consent, each runner was provided with a $120.00 voucher to purchase a pair of running shoes from a local athletic shoe store. Each participant made an appointment with the storeowner, who was an experienced runner, and upon arriving at the store underwent a treadmill running analysis performed by the owner to determine the best running shoe for the individual. Ten of the female runners selected a Brooks Ravenna running shoe, one female selected Saucony Ride running shoe and the four males selected an Asics GT 2140 running shoe. All three types of running shoes selected by the runners had a standard ethylene vinyl acetate (EVA) midsole or a biodegradable midsole with material properties very similar to EVA. All participants were instructed not to use their new running shoes until they returned to see the primary investigators.

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Instrumentation To assess in-shoe plantar pressure and vertical force data, a Novel PEDAR® capacitance sensor insole (Novel USA, Minneapolis, MN) was used with a sampling rate of 50 Hz. Hurkmanns, et al has reported good repeatability with the PEDAR® insoles when used to measure vertical force and pressure between multiple days of measurement.13 The PEDAR® capacitance sensor insole consisted of a matrix of 90 to 100 capacitance transducers and was approximately 2 mm in thickness. All six pairs of sensor insoles used in the study were calibrated prior to the start of data collection using a rubber bladder that was pressurized with compressed air over a range beginning with 5 kPa and ending with 600 kPa. The cables from the sensor insoles were attached to a Bluetooth enabled device worn by the runner that transmitted data to the computer. Thus, the participant was not tethered to the computer with cabling when running over the 42-meter runway. To assess the hardness of the midsole material, a Shore type A durometer (Model 1500; Rex Gauge Company, Inc., Buffalo Grove, IL) was utilized. The measurement obtained using a Shore type A durometer results in a value between 0 and 100 with higher values indicating a harder material. Procedures Upon arriving for the initial data collection session, each participants body weight and height were obtained. They were then asked to don their new running shoes and walk for one mile followed by a two-mile run for two consecutive days to “break-in” the new shoes. Once the shoe “break-in” period was completed, the participant returned to complete the initial FCAT as well as measurement of midsole hardness and in-shoe pressure assessment. As previously noted, Mundermann, et al reported that the FCAT had excellent levels of reliability for all nine scales if a control condition was used before each session the FCAT was utilized.11 For this study, the control shoe condition was a martial arts shoe with soft leather upper and a hard, flat rubber outsole (Tiger Claw Martial Artist’s Athletic Shoes, Pioneer Interstate, Inc., Nashville, TN). The martial arts shoe was selected as the control shoe condition instead of a standard running based on Mundermann, et al who reported that the control shoe condition should not

have a similar density in comparison to the actual shoe density being tested.11 The original cushioned sock liner (insole) from the martial arts shoe was removed and replaced with a piece of non-molded Aliplast 10 material with a thickness of three millimeters (Alimed, Inc, Dedham, MA). Aliplast 10 is a firm, closed cell polyurethane material with durometer of 58 (Shore A gauge). The participant was then asked to run at a self-selected speed used for a typical training run over a 42-meter indoor runway three times. Once they completed the practice runs in the control shoe, they were asked to don their running shoes and repeat running the same distance at the same self-selected running speed. Running speed was monitored using two infra-red photocells position 20 meters apart (Tandy Corp., Fort Worth Texas) and connected to a digital timer (model 54030; Lafayette Instrument Co, Lafayette, Indiana). To control for possible variations in the cushioned sock liner (insole) of the running shoe and to only assess the cushioning provided by the running shoe, prior to beginning their run the insole was removed from the running shoe and replaced with a piece of 3 mm non-molded Aliplast 10 material to ensure the inner aspect of the shoe was standardized for all shoes. When the participant completed running the three practice trials over the 42-meter runway with their running shoes, the participant was asked to immediately sit down and complete the FCAT after receiving verbal instructions by the same investigator (MWC). While they were completing the FCAT, another investigator (TGM) assessed the hardness in the heel region of the running shoe using the Shore type A durometer. The measurement was made within the shoe directly on top of the midsole at the center of the heel and three centimeters from the most posterior aspect of the shoe with the Aliplast 10 material removed (see Figure 1). The average of three consecutive durometer measurements was recorded. After the durometer measurement was completed, the Aliplast 10 material was placed back in the shoe and an appropriately sized PEDAR capacitance sensor insole was placed on the top of the Aliplast 10 material. The participant was then asked to run at the same selfselected pace over a 42-meter indoor runway while capacitance sensor data were collected and running speed was monitored as previously described. For all capacitance sensor data collection, the subject was

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Figure 1. Measurement of midsole hardness in the heel region of the running shoes using the durometer.

asked to not wear socks to prevent any possible cushioning effects of the sock on the capacitance sensor data. When questioned about not wearing socks, each participant indicated they did not believe that his or her gait pattern had changed when running without socks and this was confirmed by observation of each runnerâ&#x20AC;&#x2122;s gait pattern by the investigators during data collection. Following the completion of the capacitance sensor data collection, each participant was instructed to follow their usual running regime using study footwear and was provided with a calendar to log the distance run per day. Participants were instructed to notify the investigators after they had completed running 160, 320, 480, and 640 km so they could return and have all the measures described above repeated. To ensure that the participant ran at the same speed for baseline and all follow-up assessments, the maximum variation in the self-selected running speed allowed was less than 5%. Data Analysis. The FCAT ratings for each of the nine scales provided by the participants at the initial assessment and after running 160, 320, 480, and 640 kilometers were measured in mm from the zero point to determine the value of each scale related

shoe comfort, fit and cushioning. To determine the change in plantar pressure and vertical forces, ten foot strikes for the right foot only were randomly selected from the middle 15 meters of the 42-meter run. This was done to ensure that participants were neither accelerating nor decelerating from their self-selected running speed during data collection. The selection of ten foot strikes for the analysis of force and pressure data was based on the findings of Kernozek, et al, who reported high levels of reliability across all foot regions with at least nine steps while using the PEDAR insoles.14 Once the ten foot strikes were selected for each participant, the Percent Mask program (Novel USA, Minneapolis, MN) was used to divide each step into the following six plantar regions: medial rearfoot, lateral rearfoot, medial midfoot, lateral midfoot, medial forefoot, and lateral forefoot. These plantar regions were determined based on a percentage of total foot length and width and were consistently applied to all ten steps selected for analysis. The heel region was from 0% to 30%, the midfoot region from 30% to 60%, and the forefoot from 60% to 85% of total foot length. The total widths of the rearfoot, midfoot, and forefoot were divided in half. Once all the regions were defined, the Multimask Evaluation program (Novel USA, Minneapolis, MN) was used to calculate the pressure-time integral (PTI) and force-time integral (FTI) for the medial and lateral heel regions. The PTI and FTI were also selected for analysis since these two variables assess the magnitude of plantar pressure and vertical force applied during the time the medial and lateral heel regions are in contact with the supporting surface. The values for the PTI and FTI for both heel regions for all ten foot strikes were averaged and used for further statistical analysis. Statistical Analysis In addition to descriptive statistics, a series of independent t-tests were used to determine differences existed in demographics between the female and male runners. Repeated measures analysis of variance (ANOVA) tests were performed to determine if the FCAT, FTI, PTI, or heel midsole hardness changed because of running mileage. Post hoc comparisons were performed to determine differences among the test conditions. An alpha level of 0.05 was used for all tests of statistical significance.

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RESULTS The demographic data for all 15 subjects is provided in Table 1. All 15 runners were able to complete the required 640 km of running with the study footwear within 32 weeks from the first day of testing. The results of the t-tests indicate that there were no significant differences between the female and male subjects for age, body weight, or BMI. The only demographic variable that was significantly different was height. Based on these results, all statistical analysis was performed on all 15 subjects regardless of gender.

Self-Perceived Shoe Comfort, Fit, and Comfort Footwear Comfort Assessment Tool (FCAT) scores for each of the nine variables measuring self-perceived shoe comfort, fit, and comfort decreased over the course of running 640 km. These decreases ranged from 2.7% for Heel Cushioning to 10.4% for Pronation-Supination Control (Table 2). Despite these slight self-perceived reductions in perceived shoe comfort, fit, and cushioning over time, based on the repeated measures ANOVA none of the nine variables were found to be statistically significant.

Table 1. Demographic information on the subjects used in this study. Values in parentheses are standard deviations

Midsole Hardness The durometer values indicating the hardness of the midsole in the heel region are shown in Table 3. Between baseline and 640 km, heel midsole hardness increased approximately 17%. The results of the repeated measures ANOVA for the heel region durometer values were significantly different (p=.000) as running mileage increased. The results of post hoc tests indicated that the increase heel hardness was significantly different from baseline to: 160 km (p=.001), 320 km (p=.001), 480 km (p=.008) and 640 km (p=.000).

Table 2. Subject perceived shoe comfort and ďŹ t while running up to 640 kilometers as assessed using the Footwear Comfort Assessment Tool (FCAT). The values listed for each comfort variable are based on a 0 to 100-millimeter visual analogue scale with standard deviation in parentheses.

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Plantar Pressure and Vertical Force PTI values decreased approximately 20% in the medial heel and 28% in the lateral heel as the number of miles ran increased. The results of the repeated measures ANOVA for PTI were significantly different for both the medial (p=.009) and lateral heel (p=.000) as running mileage increased. Post hoc tests indicated significant differences (p<.05) between baseline and 320 km for the lateral heel compared to 480 kilometers for the medial heel

(Table 4). Like the findings for PTI, FTI values also decreased approximately 16% in the medial heel and 33% in the lateral heel as the number of miles ran increased. The results of the repeated measures ANOVA for FTI were significantly different for both the medial (p=.003) and lateral heel (p=.000) as running mileage increased. Post hoc tests indicated significant differences (p<.05) between baseline and 320 km for the lateral heel and between baseline and 480 km for the medial heel (Table 4). DISCUSSION Previous research has not assessed the ability of the runner to self-perceive changes in heel cushioning provided by the midsole when using the same pair of running shoes over an extended period of distance running. Thus, one purpose of this study was to determine the ability of the runner to self-report changes in the heel cushioning properties of the midsole in a new pair of running shoes at baseline and after running 160, 320, 480, and 640 km. These mileage distances were selected based on previous research demonstrating that typical midsole degradation in the heel region occurs prior to or at 640 km. To assess changes in the midsole cushioning properties, the PTI and FTI for the medial and lateral heel regions were assessed using an in-shoe capacitance sensor insole. Between baseline and after running 480 km, both the PTI and FTI showed reductions in the amount of midsole heel cushioning of approximately 16% to 33% with no further significant reductions after running 640 km. As would be expected

Table 3. Change in durometer values of the midsole over time. Durometer measurements were obtained using a Shore type A durometer (Model 1500; Rex Gauge Company, Inc., Buffalo Grove, IL) with a higher value indicating a harder material

Table 4. Mean plantar pressure and force values under the heel at baseline and every 160 kilometers run. The values in parentheses are standard deviations km

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since most runners with a rearfoot strike pattern tend to load the lateral aspect of the heel region of the midsole when initially contacting the supporting surface, the lateral heel region demonstrated a more rapid reduction in heel cushioning for both the PTI and FTI, than the medial heel region. Significant changes in midsole heel cushioning, based on both PTI and FTI data, were observed between baseline and 320 km for the lateral heel region but not until 480 km for the medial heel region. While these reductions in the PTI and FTI, which represent the degree of midsole degradation in the heel region, are very similar to values previously reported in the literature,6,8 none of the 15 runners self-reported a significant reduction in midsole cushioning after running 640 km. In fact, the percent decrease noted by the runners in this study using the FCAT was the least for heel cushioning (2.7%) in comparison to the other eight scales assessing comfort, fit and forefoot cushioning. Based on these findings, the first hypothesis that stated individuals who run at least 24 km (15 miles) per week would be able to perceive a reduction in the midsole cushioning when running in the same pair of shoes over a running distance of 640 km was rejected. In addition, the third hypothesis which stated that reductions in force and plantar pressures in the medial and lateral heel region of the midsole would not be greater than 20% after running 640 km was also rejected since reductions in the PTI and FTI were 28% and 33% respectively after running 480 km. The runnersâ&#x20AC;&#x2122; inability to self-perceive a loss in the level of cushioning provided by the heel region of the midsole could be attributed to the increased thickness of the midsole in the heel region. While the thickness of the midsole in the forefoot of the running shoes used in this study was 1.6 centimeters, the heel region thickness was 2.8 centimeters. While recent researchers have attributed the increased thickness of the heel region of the midsole as a factor that facilitates a rearfoot strike pattern during running,1 Robbins, et al were the first to suggest that the increased cushioning provided by running shoes can act to attenuate the perceived magnitude of forces acting on the plantar surface of the foot.15 The findings of the current study would appear to support the theory proposed by Robbins and colleagues.15 Since it would appear that the runner may not be able to self-assess changes in the cushioning properties of

the heel region of the midsole, the ability to quickly and objectively assess changes in the hardness of the material in the heel region of the midsole would be of value to the clinician. In the current study between baseline and after running 640 km, heel midsole hardness increased approximately 17% based on durometer measurements, which is similar to the decreases noted in both the PTI and the FTI. Based on these findings, the second hypothesis which stated that the use of a durometer could be used to objectively measure changes in midsole material hardness in the heel region of running shoe midsole used by the same individual over a running distance of 640 km was not rejected. The durometer used in this study was easy to use and with a cost of $250.00 is feasible in those clinical settings that specialize in the evaluation and management of running injuries. In the current study, the running footwear for each participant was assessed using the durometer prior to using the shoes which provided a baseline value for comparison. While the authors recognize that it is not always practicable in the typical clinical setting, for those clinicians who provide services or consult with high school or collegiate athletic teams or recreational running clubs, baseline durometer measurements of the midsole can be made during pre-running season screenings that would allow the clinician assess changes in midsole material hardness at a later date. Future studies are also needed to provide baseline durometer readings for various types of midsole materials. It is unclear whether a reduction in the ability of the shoe midsole to absorb forces or plantar pressures is a factor in the development of running injuries. Based on the results of their prospective study, Taunton, et al have suggested that as the running shoe ages with use, running injuries can increase as the cushioning and support qualities of the shoe decline.16 More recently, Kong, et al demonstrated that as shoe cushioning decreases, individuals modify their running patterns to maintain constant external loads and that the adaptation strategies due to shoe degradation were not affected by different cushioning technologies.7 Although the development of running injuries is multifactorial, irrespective of the specific influence of the running shoe on the development of runningrelated injuries, current evidence would support the need for the clinician to assess midsole cushioning as part of the physical examination of the runner.

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Limitations of the current study include the limited number of runners recruited to participate and that all runners in the study utilized a rearfoot strike pattern. The restriction in the number of runners who took part in the study was directly due to the cost of providing new footwear for all participants. Although only 15 runners participated, the ability to follow the change in midsole hardness, vertical forces and plantar pressures from baseline or zero to 640 km in a new pair of running shoes provides important information on self-perception and midsole durability. While only runners with a rearfoot strike pattern were selected for participation, the focus of this study was to assess the degree of cushioning degradation in the heel region of the midsole of the running shoes. The fact that previous studies have shown that the rearfoot strike pattern is the most common amongst recreational and collegiate cross-country runners would justify the selection of individuals that use a rearfoot strike pattern when running.17,18 CONCLUSION The results of this study indicate that runners who have a rearfoot strike pattern, when using a new pair of running shoes with an EVA or a biodegradable midsole with material properties, will have a 16% to 33% reduction in the amount of midsole heel cushioning after running 480 km. The degree of reduction will be greater in the lateral heel region of the midsole in comparison to the medial heel region. It would appear that even though these reductions in heel cushioning are significant, experienced recreational runners are not able to self-perceive these changes in cushioning after running 640 km. In addition, based on the findings of this study the clinician can utilize a durometer to quickly and accurately assess changes in the hardness of the heel region of the midsole as running mileage increases. REFERENCES 1. Lieberman DE, Venkadesan M, Werbal WA, et al. Foot strike patterns and collision forces in habitually barefoot versus shod runners. Nature. 2010; 463:531-535. 2. Chambon N, Sevrez V, Ly QH, et al. Aging of running shoes and its effect on mechanical and biomechanical variables: implications for runners. J Sports Sci. 2014; 32:1013-1022. 3. Cavanagh PR, Lafortune MA: Ground reaction forces in distance running. J Biomech. 1980; 13:397-406.

4. Taunton JE, McKenzie DC, Clement DB: The role of biomechanics in the epidemiology of injuries. Sports Med. 1988; 6:107-120. 5. Cook SD, Kester MA, Brunet ME: Shock absorption characteristics of running shoes. Am J Sports Med. 1985; 13:248-253. 6. Hamill J, Bates BT: A kinetic evaluation of the effects of in vivo loading on running shoes. J Orthop Sports Phys Ther. 1988; 10:47-53. 7. Kong PW, Candelaria NG, Smith DR: Running in new and worn shoes: a comparison of three types of cushioning footwear. Br J Sports Med. 2009; 43:745-749. 8. Schwanitz S, Odenwald S. Long term cushioning properties of running shoes. In: Estivalet M & Brisson P. The Engineering of Sport 7, Vol 2. 7th ed. Paris, France: Springer-Verlag; 2008:95-100. 9. Wang L, Li JX, Hong Y, et al. Changes in heel cushioning characteristics of running shoes with running mileage. Footwear Science. 2010; 2:141-147 10. Verdejo R, Mills NJ: Heel-shoe interactions and the durability of EVA foam running-shoe midsoles. J Biomech. 2004; 37:1379-1386. 11. Mundermann A, Nigg BM, Stefanyshyn DJ, Humble N: Development of a reliable method to assess footwear comfort during running. Gait Posture. 2002; 16:38-45. 12. Barton CJ, Bonanno D, Menz HB: Development and evaluation of a tool for the assessment of footwear characteristics. J Foot Ankle Res. 2009; 2:10. 13. Hurkmanns HLP, Bussman JBJ, Benda E, et al. Accuracy and repeatability of the Pedar Mobile system in long-term vertical force measurements. Gait Posture. 2006; 23:118-125. 14. Kernozek TW, LaMott EE, Dancisak MJ: Reliability of an in-shoe pressure measurement system during treadmill walking. Foot Ankle Int. 1996; 17:204-209. 15. Robbins SE, Hanna A, Jones LA: Sensory attenuation induced by modern athletic footwear. J Testing Eval. 1988; 16:412-416. 16. Tuanton JE, Ryan MB, Clement DB, et al. A prospective study of running injuries: the Vancouver sun run â&#x20AC;&#x153;in trainingâ&#x20AC;? clinics. Br J Sports Med. 2003; 37:239-244. 17. Goss DL, Lewek M, Yu B, et al. Lower extremity biomechanics and self-reported foot-strike patterns among runners in traditional and minimalist shoes. J Ath Train. 2015;603-611. 18. Bade MB, Aaron K, McPoil TG: Accuracy of selfreported foot strike pattern in intercollegiate and recreational runners during shod running. Int J Sports Phys Ther. 2016; 11:64-71.

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IJSPT

ORIGINAL RESEARCH

THERE ARE NO BIOMECHANICAL DIFFERENCES BETWEEN RUNNERS CLASSIFIED BY THE FUNCTIONAL MOVEMENT SCREEN Rodrigo Ribeiro de Oliveira¹,² Shalimá Figueirêdo Chaves¹ Yuri Lopes Lima¹ Márcio Almeida Bezerra¹,² Gabriel Peixoto Leão Almeida¹,² Pedro Olavo de Paula Lima¹,²

ABSTRACT Background: Running has been one of the main choices of physical activity in people seeking an active lifestyle. The Functional Movement Screen (FMS™) is a screening tool that aims to discern movement competency. Purpose: The purposes of this study were to compare biomechanical characteristics between two groups rated using the composite FMS™ score, and to analyze the influence of specific individual tests. The hypothesis was that the group that scored above 14 would demonstrate better performance on biomechanical tests than the group that scored below 14. Study Design: Cross-Sectional Study. Methods: Runners were screened using the FMS™ and were dichotomized into groups based on final score: Functional, where the subjects scored a 14 or greater (G≥14, n=16) and dysfunctional, when the subjects scored less than 14 (G< 14, n=16). All runners were evaluated using measures for flexibility, postural balance, muscle strength, knee dynamic valgus during forward step down test and time for the electromyographic response of the transversus abdominis and fibularis longus muscles. All data were analyzed with SPSS (p≤0.05) and the index of asymmetry (IS) was calculated with the mean score of nondominant limb divided by the mean score of the dominant limb, multiplied by 100. Results: There were no statistically significant differences in flexibility, muscle strength, knee dynamic valgus, or myoelectric response time of the transversus abdominis and long fibular muscles. Index of asymmetry (IS) of global stability was 3.26±26.79% in G≥14 and 31.72±52.69% in G<14 (p=0.02). In-line lunge and active straight-leg raise tests showed no significant difference between the groups (p>0.05). Conclusions: Overall, there were no biomechanical differences between the groups of runners as classified by the FMS™. In addition, in-line lunge and active strength-leg raise tests did not influence on the FMS™ final score. Level of Evidence: 2b Key words: Electromyography, fundamental movements, running

League of Sports Physical Therapy, Federal University of Ceará, Brazil 2 Department of Physical Therapy, Federal University of Ceará, Brazil

CORRESPONDING AUTHOR Pedro Olavo de Paula Lima Analysis Laboratory of Human Movement Address: Rua Alexandre Baraúna, 949, Fortaleza, CE, 60430-110, Brazil Phone: +55 85 33668632 Facsimile: +55 85 33668002 E–mail: pedrolima@ufc.br

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INTRODUCTION Running is a popular sport in world because it is associated with a healthy life style, reduction of cardiovascular risk factors, and is easy to do and can be done anywhere. However, a runner needs to have knowledge about running-related injuries and risk factors for injury in order to participate in this sport safely. The incidence rates of injury in runners ranged 2.5 to 12.1 injuries per 1000 hours of running. The highest incidence running-related musculoskeletal injuries among runners are patellar tendinopathy (5.5%22.7%), medial tibial stress syndrome (9.1%-19.0%), Achilles tendinopathy (13.6%-20.0%), plantar fasciitis (4.5%-10%), patellofemoral syndrome (5.5%), and iliotibial band syndrome (1.8%-9.1%).1 Researchers indicate that factors such as deficits in postural balance, flexibility, hamstring/quadriceps ratio and activation of stabilizing muscles of the lumbar spine (such as the transversus abdominis), hip stabilizers (i.e., gluteus medius) and ankle stabilizers (i.e., fibularis longus), are related to the high incidence of these injuries.1-6 Good performance on functional assessment tools or screens has been associated with lower risk of injury in runners.7-9 The Functional Movement Screen (FMS™) is one such screening system used to evaluate the dynamic capacity of individuals in specific movements that require balance, mobility and stability, comparing the performance of the runners above and below of 14 cut-off score. The FMS™ consists of seven fundamental movements that are graded from 0 to 3 according to the performance in the execution of each movement; where a score of three means satisfactory movement competency, a score of two means that the person is able to complete the movement but with compensation, a score of one means that the person is unable to complete the movement pattern and a score of zero if at any time during the testing the person has pain. According to the final score, an individual’s results can be dichotomized as satisfactory movement competency (i.e., scores above 14) or unsatisfactory movement competency (i.e., scores below 14). According to these principles, it is assumed that individuals with optimal functional movement patterns might exhibit symmetry in variables such as strength and posture stability, good measures of flexibility, and effective activation of stabilizing muscles. In contrast, individuals with dysfunctional movement patterns might exhibit

asymmetries in variables, decreased flexibility, and difficulty with effective recruitment of muscles used to stabilize the body.10-13 The purposes of this study were to compare biomechanical characteristics between two groups of runners rated using the composite FMS™ score, and to analyze the influence of specific individual tests. The hypothesis was that the group that scored above 14 would demonstrate better performance on biomechanical tests than the group that scored below 14. METHODS A cross-sectional study was conducted with 32 runners in the Human Motion Analysis Laboratory of Department of Physical Therapy at the Federal University of Ceará. In order to be included, runners had to maintain a weekly workout routine of a minimum volume of 20 km per week and a minimum frequency of twice per week. Participants had to be between 18 and 60 years of age, and they could not have any diseases of the cardiorespiratory system, such as uncontrolled hypertension, angina, or have acute musculoskeletal pain. All subjects signed the informed consent and submitted to an interview, to identify their age, gender, weight, height, sports practice time, training volume and presence of injury over the prior year. This study was approved by the Institutional Human Research Ethics Committee (protocol number #208.176) OUTCOME MEASURES Functional Movement Screen (FMS™) The FMS™ is a screening tool used to analyze fundamental patterns of movement. It consists of seven basic movements that require balance, mobility and stability: Deep squat, hurdle step, in-line lunge, shoulder mobility, active straight leg raise (ASLR), trunk stability push-up, rotatory stability. The assessment followed the order described by the authors of the method, and each movement was scored according to the criteria described by them, where each activity could be attempted three times, and were graded from 0 to 3. Zero (0) indicated pain during execution, one (1) indicated that the individual was unable to perform the movement, two (2) suggested the individual was able to perform the movement with some compensation, and three (3) suggested the subject was capable to perform the full movement without

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any compensations. At the end, scores were summed in order to obtain the total score (composite score), which has a maximum of 21 points. If there was a different score on each limb for a bilateral test, the lower of the two scores was used in the composite score. Individuals with total score lower than 14 points were placed in a group defined by the authors as functional (Group - G<14), and those with total score of 14 or higher were placed in a group defined by the authors as dysfunctional (Group - G≥14).14,15 The participants were assessed by the same examiner, who was blinded from performance on the other tests. Flexibility Flexibility was evaluated with the sit and reach test.16 This is performed with the patient seated, with hips flexed, knees extended and foot touching the anterior surface of the bench. The participants were instructed to move the bar as far as possible by flexing the trunk and keep the position for three seconds. The distance was measured in centimeters. Three attempts were performed and the highest value was considered for analysis.17 Postural Balance Postural balance was evaluated using Biodex Balance System® (BBS®).18 This device measures the degree of sway on two axes (anterior/posterior and medial/ lateral) during testing.19 The displacement of the center of gravity in both anterior/posterior and medial/ lateral directions were analyzed as well as the overall displacement, which is a measurement obtained considering both. The protocol required one-legged stance and the athletes were positioned on the platform with tested the knee held at 10⬚ of flexion and contralateral knee flexed to 90⬚. Athletes were instructed to remain steady during the test. This assessment was repeated three times for each limb and lasted for 20 seconds with 10 seconds of interval time, and average was utilized.20 The Overall Stability Index, Antero-Posterior Stability Index, and Latero-Medial Stability Index (as calculated and provided by the BSS) were assessed and considered as outcome measures. Myoelectric response time of the TrA/internal oblique (TrA/IO) muscles Surface electromyography Miotool 400 (Miotec®, Porto Alegre/RS, Brazil) was used to evaluate the

difference in activation time between the fibers of the anterior fibers deltoid and the TrA when a rapid flexion motion of the shoulder was conducted (anticipatory contraction mechanism). Skin preparation and electrode placement were conducted following the recommendations of the SENIAM (Surface Electromyography for the Non-Invasive Assessment of Muscles).21 A pair of electrodes was placed in a horizontal position: 20 mm medially and inferiorly the anterior superior iliac spine (ASIS) to evaluate the TrA muscle, and other pair was placed to two inches below and forward from the acromion, to evaluate the deltoid muscle. The distance between the centers of the electrodes was 20 mm, and the side evaluated was always the dominant one. The reference electrode was placed on the elbow of the dominant arm.22,23 Athletes were verbally asked to flex the shoulder three times as quickly as possible to 90⬚.24 The participants were allowed to execute two to five repetitions of training in predetermined distances and speeds to familiarize them with the movement. The task was considered appropriate when the individual was able to contract the TrA muscle before or at the same time as the deltoid muscle.24 Electromyographic signals were sampled at a frequency of 2000 Hz and filtered with a band-pass range of 20-450 Hz. Myoelectric response time of ﬁbularis longus muscle In this test, the same surface electromyography unit (Miotec®, Porto Alegre/RS, Brazil) was used to identify the reaction time of the fibularis longus muscle following an inversion perturbation. If the reaction time is higher than 90 ms, there is a greater risk for ankle sprain.25 The standards of the SENIAM were followed as in the previous electromyographic examination. The pairs of electrodes were placed in the upper third of the distance between the head of the fibula and the lateral malleolus (fibularis longus muscle) and on the fifth metatarsal (for vibration assessment only). The reference electrode was placed in the lateral epicondyle of the humerus. The assessment was conducted using the dominant limb.21 The athlete was positioned in one-legged stance on a balance board, and a 10 kg weight plate dropped on the lateral surface of the board, creating a sudden disturbance, and causing ankle inversion of approximately 15⬚. The athlete was not warned

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when the weight plate would drop. The balance board was aligned with the slit, where the weight was released, in order to ensure that the weight plate would fall in the same place (Figure 1).25 Note: Video Clip 1, available online, shows how this test was administered. Electromyographic signals were sampled at a frequency of 2000 Hz and filtered with a band-pass range of 20-450 Hz. Any activities higher than three standard deviations values added to average value for the rest root mean square (RMS) was considered a reactive muscle contraction. The signal generated by the vibration of the 5th metatarsal was considered the instant which the highest frequency was perceived.26 Dynamic valgus The drop-jump vertical test was used to assess the dynamic valgus occurring at the knee.27 Athletes were instructed to drop from a 50 cm tall box, and asked to immediately execute a two-legged maximal vertical jump after dropping. Each subject performed three attempts. It was not considered valid if the athlete jumped instead of dropping or if the individual

Figure 1. Setup for inversion perturbation. A 10-kg weight was dropped onto the posterolateral edge of the balance board to create a sudden inversion stress.

was clearly out of balance. The test was filmed with a digital camera (Sony Cyber-shot DSCW35; 7.2 megapixels) in the frontal plane, while supported on a tripod, whose center was three meters away from the center of the platform used. The ability to control the hip and to avoid dynamic valgus knee during the test was classified from 0 to 2. The zero (0) indicated no significant lateral tilt of the pelvis, no valgus motion of the knee and no medial/lateral side-to-side movements of knee during performance; therefore, suggesting a good performance. Individuals graded one (1) showed movements combined or in isolation: lateral tilt of pelvis, knee slightly moving into a valgus position and some medial/lateral side-to-side movements of the knee during the movement; therefore, suggesting a reduced performance. The participants graded two (2) performed the following movements combined or in isolation: lateral tilt of the pelvis, knee clearly moving into a valgus position and medial/lateral side-to-side movements of the knee, suggesting a poor performance.28 Quadriceps and Hamstring Strength A strength test was performed with an isokinetic dynamometer (Biodex System 4; Biodex Medical Systems, New York, NY, USA).29 This test aims to assess muscle torque production at a constant velocity. The results allow the assessment of the hamstrings/quadriceps ratio. Athletes warmed-up on a stationary bike for five minutes. The dynamometer chair was positioned so that the hip was flexed at 85⬚ and the machine axis was aligned with the sagittal plane axis of the knee. Then, the participants were seated in the dynamometer chair and their positions were stabilized with belts placed at the trunk level. Abdomen and thigh belts were firmly fastened, in order to prevent undesired movements. The machine’s lever arm was fixed above the medial malleolus. The test protocol consisted of concentric isokinetic assessment at two speeds: 60⬚/s and 300º/s, with 5 and 15 repetitions respectively, and an interval of 30 seconds for rest. The equipment was calibrated with range of motion starting from a maximum flexion up to a maximum extension of the knee where the reference point was 90⬚ of flexion. The testing limb was weighed at maximal extension (180⬚) by the equipment to avoid bias caused by gravity. The upper limbs were positioned laterally

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holding the chair handles. In order to complete the muscle warm-up period and to familiarize with the equipment, participants were asked to perform five knee flexion-extension submaximal repetitions after the procedures for positioning mentioned above.30 Statistical Methods Data were analyzed in the SPSS version 17.0 software with a significance level of p ≤ 0.05. The Chi-square test was used to analyze the association of nominal variables between groups, and the independent t-test was used to compare the continuous variables. Index of asymmetry (IS) was calculated for all continuous data by the following formula: IS = (nondominant limb/dominant limb) x 100. Independent t-test was used to analyze the isolated influence of each FMS™ test on the final score. RESULTS The sample was composed of 32 runners who were dichotomized into two groups: functional (n=16) and dysfunctional (n=16), according to FMS™ scores. There were no drop outs and no significant differences between groups in baseline demographic variables (Table 1). In the flexibility assessment, the functional group obtained an average of 30.71 ± 8.3 cm, while the dysfunctional group showed 25.51 ± 9.30 cm (p=0.42). In the myoelectric response time TrA assessment, three participants (18.75%) from the functional group showed responses at the right time, while in the dysfunctional group only two individuals (12.5%) were able to contract before or at the same time as the deltoid.

No difference was demonstrated between groups (␹2 = 0.654; p = 0.57). The assessment of the fibularis longus response time to a sudden inversion perturbation showed that nine participants (56.2%) from the functional group had a contraction in the proper time, while eleven participants (68.8%) from the dysfunctional group reacted in the same period of time (p=0.19). There was no difference between groups. In the drop-jump vertical test, the functional group had six individuals (40%) performing with good performance, three participants (20%) performing with reduced performance and six (40%) performing with poor performance. In the dysfunctional group, six participants (42.8%) had a good performance, four participants (28.6%) had a reduced performance, and four (28.6%) had a poor performance. There was no difference between groups (␹2=0.51; p = 0.77). Index of asymmetry was used to analyze balance. Global stability, index anterior/posterior and index medial/lateral were compared. The global stability statistically differed between groups; however, no significant differences were observed for the anterior/posterior and medial/lateral index. The results are shown in Table 2. In assessment of muscular strength, muscle strength index of asymmetry for quadriceps, hamstrings and agonist/antagonist ratio at 60º and 300⬚/s were compared between groups. No differences were detected as observed in Table 3. Data comparison of individual’ scores in FMS™ tests between groups showed that almost all tests influ-

Table 1. Characteristics of the sample Variables Age (years) Gender

Female Male

Weight (kg) Height (cm) Motor dominance

Right-footed Left-footed

Injury in last 12 months Time experience in running (years) Frequency per week Distance per week (km) a Mean ± Standard Deviation.

Functional (n=16) 43.38 ± 8.48 a 03 (9.4%) 13 (40.6%) 70.81 ± 8.96 a 168 ± 5 a 11 (34.4%) 05 (15.6%) 06 (37.5%) 7.56 ± 3.94 a 3.31 ± 0.87 a 33.19 ± 10.4 a

Dysfunctional (n=16) 39.19 ± 8.53 a 05 (15.6%) 11 (34.4%) 75.81 ± 8.36 a 171 ± 7 a 15 (46.9%) 01 (3.1%) 11 (68.7%) 5.88 ± 3.57 a 3.44 ± 1.03 a 29.81 ± 9.0 a

p 0.71 0.41 0.11 0.28 0.08 0.08 0.21 0.71 0.20

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Table 2. Comparison outcomes of balance between limbs, using the Biodex Balance System Functional Dysfunctional (n=16) (n=16) Global stability 3.26±26.79 % 31.72±52.69 % Index APa 12.74±51.76 % 51.76±98.10 % Index MLb 5.9±20.64 % 20.64±66.33 % a AP: Antero/Posterior; bML: Medial/lateral. †Significant at p < 0.05 level. Variables

p 0.02† 0.26 0.50

Table 3. Results comparing index of symmetry of isokinetic strength between limbs and agonist/antagonist values to the suggested reference standards Functional (n=16) 11.56±10.96 % 14.20±17.37 % 0.70±8.88 % 0.20±10.09 % 1.61±36.12 % 2.10±13.05 %

Variables Quadriceps 60º/s Hamstring 60º/s Quadriceps 300º/s Hamstring 300º/s Hamstring / Quadriceps ratio 60º/s Hamstring / Quadriceps ratio 300º/s

enced in the subject’s classification, with the exception of in line lunge and active straight-leg raise test, which showed no significant difference between the groups. The data are shown in Table 4. DISCUSSION The findings of this study showed that the groups classified by the total FMS™ score as either functional or dysfunctional did not demonstrate significant differences in the observed biomechanical tests or anthropometric characteristics, and the inline-lunge and active-strength-leg-raise tests did not influence the FMS™ categorization of this sample. Sample characterization data, such as age, gender, weight, height and motor dominance were compared between groups and there were no significant differences. There was no association between previous

Dysfunctional (n=16) 12.15±11.7 1% 8.12±7.32 % 3.05±17.70 % 3.30±19.37 % 3.29±22.23 % 4.04±32.88 %

p 0.88 0.20 0.45 0.52 0.87 0.50

injuries and total FMS™ score. Although no statistically significant difference existed between groups regarding previous injuries, it is important to mention that previous injury approached a statistically significant difference (p=0.08), and this relationship might be significant with a bigger sample size. Further research with bigger sample sizes is necessary to check the relation of these data. Data from training, years of experience in running, weekly frequency, and training volume were not influential factors in the classification of the sample. Despite these findings, some authors have shown an association between increased distance traveled during training and high risk of injury to the lower extremities, as well as lower volume of practice time, previous injuries, duration of each training session, and workout speed. These last two variables were not analyzed in this study.31-34

Table 4. Comparison between the groups regarding the individual scores of the FMS™ tests Tests Deep Squat Hurdle Step In line Lunge Shoulder Mobility Active Straight-Leg Raise Trunk Stability Push-up Rotary Stability † Not significant at 5% level.

Functional (n=16) 2.06±0.44 2.25±0.44 2.31±0.47 2.63±0.50 2.25±0.85 2.44±0.81 1.70±0.48

Dysfunctional (n=16) 1.56±0.51 1.88±0.34 2.06±0.57 1.44±0.90 1.94±0.68 1.31±0.80 1.31±0.48

p 0.006 0.01 0.20† 0.001 0.26† 0.001 0.03

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The sit and reach test for flexibility was not different between groups. However, this test of the posterior chain should in theory be similar to the result of the ASLR test, which also evaluates active posterior chain flexibility and thus should have been different between groups. The authors believe that this occurred because one test involves the spine in the procedure and the other does not. The findings of the current study are not in agreement with the findings of a study of 64 US Army soldiers, which showed that the passive flexibility measured by an inclinometer is strongly associated with good performance in FMS™. This discrepancy may be explained because one group was comprised of runners and the other one was comprised of soldiers, and they may have different physical requirements for performance of their chosen activities.14,35 The transversus abdominis activation data also showed no association with the classification of functional assessment, agreeing with the findings of Okada, Huxel, Nesser36 who examined 28 healthy subjects who aimed to determine the relationship between low back stability, FMS™ and performance. The results of the current study demonstrated that abdominal muscular activation was not correlated with functional performance as measured by the FMS™, despite it being present in many of the prevention programs designed to prevent musculoskeletal injuries.36 The authors think that this occurred because the FMS™ does not examine rapid movements, nor does it attempt to predict runners susceptible to low back pain. The fibularis longus muscle response time was not associated with the classification of FMS™, even though it has been described it is a risk factor for ankle sprains.37 The authors think that this occurred because the FMS™ does not examine rapid movements or factors related to muscular activation, nor does it attempt to predict athletes susceptible to ankle sprains. Similar findings occurred with test for dynamic valgus, which showed no difference between the groups in the sample, highlighting again that the FMS™ does not test this type movement. This does not negate the importance of dynamic valgus as a risk factor for development of patellofemoral syndrome or anterior knee pain, which is a common injury among runners,38 and the results

highlight why tests for dynamic valgus should be included in athlete screening. Body balance was not an important factor between the groups; however, the global stability statistically differed between groups. This could be because of how balance was measured, as that the units (degrees) from the BBS are unique to this testing device and the global measure takes into account both other measures. The current results corroborate with the findings of another study that evaluated dynamic balance in the military where no difference in a measure of balance was found between groups, though the Y-balance test was performed in the aforementioned study instead of BBS®.35 Although the results of these two studies agree (no difference in balance between groups) the Y-Balance test is not correlated with the BBS®.39 The analyzed muscle strength variables showed no differences between groups for body symmetry, indicating that both groups were symmetrical in the measured variables. This finding matches with the findings of a study that used the isokinetic test to dynamically evaluate the peak of torque, motor dominance and balance between agonists and antagonists of runners and compared with a non-athlete population. They concluded that healthy runners have the characteristic symmetry of muscular strength of knee flexors and extensors.40 Considering the isolated tests that comprise the FMS™, the in-line lunge and active straight-leg raise had no differences between the groups, suggesting that these two tests do not influence the classification of subjects into functional or dysfunctional groups using total FMS™ scores. A study was conducted comparing the performance of in-line lunge test with plantar pressure distribution, maximum jump height and a 36 m race time in 35 active and healthy subjects. No correlation with these tests was observed, suggesting that the good performance on in line lunge might not have a relation with better functional performance.41 Some authors have suggested that the cutoff point (a score of 14) used as standard for FMS™ in some populations may not be suitable for runners and other athletes. The hypothesis of this research was to state that there would be a difference in the

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results of biomechanical tests between groups, and that individuals classified as good quality movement (as defined by a score > 14) would perform better than individuals who were classified as poor quality (dysfunctional) movement (as defined as a score) lower than 14. The findings of this research showed that no group exhibited better performance on a variety of biomechanical tests, suggesting that the FMS™ may not be a suitable test for assessing risk of injury for this type of athlete, or this cut point may not be suitable for this population as has been suggested by recent studies. This is an important point that requires more detailed investigation.8,9 The authors believe that the results found may be due to a small sample size, as well as the incapacity of FMS™ to predict injuries due to the cutoff point not being appropriate for the population assessed in the current study, as the value was originally described for football players. The cutoff point is currently the most contested explanation, and considered a key point for the authors, requiring additional study.42,43 CONCLUSION The results of the current study demonstrated no differences in biomechanical measures between the groups classified as functional or dysfunctional by the total FMS™ scores. This may indicate that the cut off total FMS™ score of 14 may not properly categorize runners. In addition, the in-line lunge and active straight-leg raise tests did not directly influence the classification of individuals in subgroups. REFERENCES 1. Lopes AD, Hespanhol Junior LC, Yeung SS, et al. What are the main running-related musculoskeletal injuries? A Systematic Review. Sports Med. 2012;42(10):891-905. 2. Hrysomallis C. Relationship between balance ability, training and sports injury risk. Sports Med. 2007;37(6):547-556. 3. Hides JA, Stanton WR, Mendis MD, et al. Effect of motor control training on muscle size and football games missed from injury. Med Sci Sports Exerc. 2012;44(6):1141-1149. 4. Childs JD, Teyhen DS, Casey PR, et al. Effects of traditional sit-up training versus core stabilization exercises on short-term musculoskeletal injuries in US Army soldiers: a cluster randomized trial. Phys Ther. 2010;90(10):1404-1412.

5. Felicio LR, Baffa AP, Liporacci RF, et al. Analysis of patellar stabilizers muscles and patellar kinematics in anterior knee pain subjects. J Electromyogr Kinesiol. 2011;21(1):148-153. 6. Reeser JC, Verhagen E, Briner WW, et al. Strategies for the prevention of volleyball related injuries. Br J Sports Med. 2006;40(7):594-600; discussion 599-600. 7. Hotta T, Nishiguchi S, Fukutani N, et al. Functional Movement Screen for Predicting Running Injuries in 18- to 24-Year-Old Competitive Male Runners. J Strength Cond Res. 2015;29(10):2808-2815. 8. Loudon JK, Parkerson-Mitchell AJ, Hildebrand LD, et al. Functional movement screen scores in a group of running athletes. J Strength Cond Res. 2014;28(4):909-913. 9. Agresta C, Slobodinsky M, Tucker C. Functional Movement ScreenTM - Normative Values in Healthy Distance Runners. Int J Sports Med. 2014. 10. Cook G, Burton L, Hoogenboom BJ, et al. Functional movement screening: the use of fundamental movements as an assessment of function-part 2. Int J Sports Phys Ther. 2014;9(4):549-563. 11. Cook G, Burton L, Hoogenboom BJ, et al. Functional movement screening: the use of fundamental movements as an assessment of function - part 1. Int J Sports Phys Ther. 2014;9(3):396-409. 12. Cuchna JW, Hoch MC, Hoch JM. The interrater and intrarater reliability of the functional movement screen: A systematic review with meta-analysis. Phys Ther Sport. 2016;19:57-65. 13. Parenteau-G E, Gaudreault N, Chambers S, et al. Functional movement screen test: A reliable screening test for young elite ice hockey players. Phys Ther Sport. 2014;15(3):169-175. 14. Cook G, Burton L, Hoogenboom B. Pre-participation screening: the use of fundamental movements as an assessment of function - part 1. N Am J Sports Phys Ther. 2006;1(2):62-72. 15. Cook G, Burton L, Hoogenboom B. Pre-participation screening: the use of fundamental movements as an assessment of function - part 2. N Am J Sports Phys Ther. 2006;1(3):132-139. 16. Ayala F, Sainz de Baranda P, De Ste Croix M, et al. Absolute reliability of five clinical tests for assessing hamstring flexibility in professional futsal players. J Sci Med Sport. 2012;15(2):142-147. 17. Bertolla F, Baroni BM, Leal Junior EC, et al. Effects of a training program using the Pilates method in flexibility of sub-20 indoor soccer athletes. Rev Bras Med Esporte. 2007;13(4):222-226. 18. Arifin N, Abu Osman NA, Wan Abas WA. Intrarater test-retest reliability of static and dynamic stability indexes measurement using the Biodex Stability

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19.

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System during unilateral stance. J Appl Biomech. 2014;30(2):300-304. Arnold BL, Schmitz RJ. Examination of balance measures produced by the biodex stability system. J Athl Train. 1998;33(4):323-327. Alonso AC, Brech GC, Bourquin AM, et al. The inﬂuence of lower-limb dominance on postural balance. Sao Paulo Medical Journal. 2011;129(6):410413. Hermens HJ, Freriks B, Disselhorst-Klug C, et al. Development of recommendations for SEMG sensors and sensor placement procedures. J Electromyogr Kinesiol. 2000;10(5):361-374. Hodges PW, Richardson CA. Inefﬁcient muscular stabilization of the lumbar spine associated with low back pain. A motor control evaluation of transversus abdominis. Spine. 1996;21(22):2640-2650.

23. Marshall P, Murphy B. The validity and reliability of surface EMG to assess the neuromuscular response of the abdominal muscles to rapid limb movement. J Electromyogr Kinesiol. 2003;13(5):477-489. 24. Hodges PW, Richardson CA. Feedforward contraction of transversus abdominis is not influenced by the direction of arm movement. Exp Brain Res. 1997;114(2):362-370. 25. Briem K, Eythorsdottir H, Magnusdottir RG, et al. Effects of kinesio tape compared with nonelastic sports tape and the untaped ankle during a sudden inversion perturbation in male athletes. J Orthop Sports Phys Ther. 2011;41(5):328-335. 26. Neptune RR, Kautz SA, Hull ML. The effect of pedaling rate on coordination in cycling. J Biomech. 1997;30(10):1051-1058. 27. Redler LH, Watling JP, Dennis ER, et al. Reliability of a field-based drop vertical jump screening test for ACL injury risk assessment. Phys Sportsmed. 2016;44(1):46-52. 28. Stensrud S, Myklebust G, Kristianslund E, et al. Correlation between two-dimensional video analysis and subjective assessment in evaluating knee control among elite female team handball players. Br J Sports Med. 2011;45(7):589-595. 29. Montgomery LC, Douglass LW, Deuster PA. Reliability of an isokinetic test of muscle strength and endurance. J Orthop Sports Phys Ther. 1989;10(8):315-322. 30. Chaves SF, Marques NP, Silva RL, et al. Neuromuscular efficiency of the vastus medialis obliquus and postural balance in professional soccer athletes after anterior cruciate ligament reconstruction. Muscles Ligaments Tendons J. 2012;2(2):121-126. 31. Hespanhol Junior LC, Costa LOP, Lopes AD. Previous injuries and some training characteristics

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predict running-related injuries in recreational runners: a prospective cohort study. J Physiother. 2013;59:263-269. Van Gent BR, Siem DD, Van Middelkoop M, et al. Incidence and determinants of lower extremity running injuries in long distance runners: a systematic review. Br J Sports Med. 2007; 41(8):46980. Poppel D, Scholten-Peeters G, Middelkoop M, et al. Prevalence, incidence and course of lower extremity injuries in runners during a 12-month follow-up period. Scand J Med Sci Sports. 2013; 24(6):943-9. Hespanhol Junior LC, Costa LO, Carvalho AC, et al. A description of training characteristics and its association with previous musculoskeletal injuries in recreational runners: a cross-sectional study. Braz J Phys Ther. 2012;16(1):46-53. Teyhen DS, Shaffer SW, Lorenson CL, et al. Clinical measures associated with dynamic balance and functional movement. J Strength Cond Res. 2014;28(5):1272-1283.

36. Okada T, Huxel KC, Nesser TW. Relationship between core stability, functional movement, and performance. J Strength Cond Res. 2011;25(1):252-261. 37. Sefton JM, Hicks-Little CA, Hubbard TJ, et al. Sensorimotor function as a predictor of chronic ankle instability. Clin Biomech (Bristol, Avon). 2009;24(5):451-458. 38. Petersen W, Ellermann A, Gosele-Koppenburg A, et al. Patellofemoral pain syndrome. Knee Surg Sports Traumatol Arthrosc. 2014;22(10):2264-2274. 39. Almeida GPL, Monteiro IS, Marizeiro DF, et al. Y balance test has no correlation with the Stability Index of the Biodex Balance System. Musculoskelet Sci Pract. 2017; 27:1-6. 40. Siqueira CM, Pelegrini FR, Fontana MF, et al. Isokinetic dynamometry of knee ﬂexors and extensors: comparative study among non-athletes, jumper athletes and runner athletes. Rev Hosp Clin Fac Med Sao Paulo. 2002;57(1):19-24. 41. Hartigan EH, Lawrence M, Bisson BM, et al. Relationship of the functional movement screen in-line lunge to power, speed, and balance measures. Sports Health. 2014;6(3):197-202. 42. Frost DM, Beach TA, Campbell TL, et al. An appraisal of the Functional Movement Screen™ grading criteria–Is the composite score sensitive to risky movement behavior? Phys Ther Sport. 2015;16(4):324-330. 43. Fox D, O’Malley E, Blake C. Normative data for the functional movement screen™ in male Gaelic ﬁeld sports. Phys Ther Sport. 2014;15(3):194-199.

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IJSPT

ORIGINAL RESEARCH

RUNNING INJURY DEVELOPMENT: THE ATTITUDES OF MIDDLE- AND LONG-DISTANCE RUNNERS AND THEIR COACHES Karen Krogh Johansen, MSc1 Adam Hulme, MSc2 Camma Damsted, MSc1 Daniel Ramskov, MSc1,3 Rasmus Oestergaard Nielsen, PhD1

ABSTRACT Background: Behavioral science methods have rarely been used in running injury research. Therefore, the attitudes amongst runners and their coaches regarding factors leading to running injuries warrants formal investigation. Purpose: To investigate the attitudes of middle- and long-distance runners able to compete in national championships and their coaches about factors associated with running injury development. Methods: A link to an online survey was distributed to middle- and long-distance runners and their coaches across 25 Danish Athletics Clubs. The main research question was: “Which factors do you believe influence the risk of running injuries?”. In response to this question, the athletes and coaches had to click “Yes” or “No” to 19 predefined factors. In addition, they had the possibility to submit a free-text response. Results: A total of 68 athletes and 19 coaches were included in the study. A majority of the athletes (76% [95%CI: 66%; 86%]) and coaches (79% [95%CI: 61%; 97%]) reported “Ignoring pain” as a risk factor for running injury. A majority of the coaches reported “Reduced muscle strength” (79% [95%CI: 61%; 97%]) and “high running distance” (74% [95%CI: 54%; 94%]) to be associated with injury, while half of the runners found “insufficient recovery between running sessions” (53% [95%CI: 47%; 71%]) important. Conclusion: Runners and their coaches emphasize ignoring pain as a factor associated with injury development. The question remains how much running, if any at all, runners having slight symptoms or mild pain, are able to tolerate before these symptoms develop into a running-related injury. Level of Evidence: 3b Keywords: Attitudes, coach, etiology, running injury

Department of Public Health, Section of Sport Science, Aarhus University, DK-8000 Aarhus 2 Australian Collaboration for Research into Injury in Sports and its Prevention (ACRISP), Federation University Australia, SMB Campus, PO Box 663, Ballarat, VIC, 3353, Australia 3 University College Northern Denmark, Department of Physiotherapy, Aalborg, Denmark Funding: None Conflicts of Interest: None of the authors had any conflicts of interest, including relevant financial interests, activities, relationships, and affiliations.

CORRESPONDING AUTHOR Rasmus Oestergaard Nielsen Department of Public Health, Section of Sport Science, Faculty of Health Science, Aarhus University Dalgas Avenue 4, room 438, DK-8000 Aarhus C. Denmark E-mail: roen@ph.au.dk

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INTRODUCTION The middle- and long-distance (M/L) running disciplines are a major part of organized athletics around the world.1-3 M/L running events range from 800-meter track races to marathon running as well as cross-country. Indeed, elite runners who have high training loads can exceed 35 hours per week of running prior to a given major championship.2,4 Intense training regimes and multiple competitive events across the sporting calendar can place a considerable level of stress on the athlete’s body. Consequently, runners competing at a high level are vulnerable to musculoskeletal injuries1-3 that can be serious enough to debilitate those afflicted.5 Across the disciplines in athletics, authors have reported an injury prevalence proportion of 43-76% among the athletes,2,6 whereas studies investigating the incidence of injuries during an entire athletics season reported that two out of three athletes sustain injuries on an annual basis.2,7 M/L runners and their coaches strive to identify interventions to reduce running-related injuries (RRI’s) despite a limited evidence-based knowledge to support their efforts.8 Since effective intervention strategies build on etiological evidence,9 the first scientific hurdle to clear is to increase the knowledge about injury etiology. Since the 1970s, scientific interest around the development of RRI’s has steadily increased and considerable efforts have been made to shed light on these injuries.10 Notwithstanding the identification of few statistically significant risk and protective factors for running injury amongst M/L runners, such as a history or previous injury,4 little overall progress about the etiology of RRI’s has been made to date.2,11 The routine application of traditional epidemiologic approaches has been, and will remain to be, a necessary step for better understanding the etiology of RRI´s. However, the use of original, complementary and alternative research approaches is also required alongside common practice if the complex origins of distance running injury are to be realized.12-14 Given that quantitative studies have largely dominated the RRI literature, there is a need for more qualitative research approaches that directly include end-users and other key members of ‘the distance running system’.12 Qualitative methodologies in the context

of healthcare research presents many advantages if deemed appropriate given the circumstances and specific research questions asked.15,16 For example, value-laden questions, complex health-related phenomena not easily reducible to their component parts, and particular research outcomes that are partly characterized by social and behavioral influences are well suited to qualitative research approaches.17 Ascertaining certain data and information from middle- and long distance runners and their coaches is essential for appreciating participants’ values, principles and motivations underpinning certain actions and behaviors.18 Not only can qualitative research elucidate runners’ attitudes and experiences about their own psychosocial dispositions preceding injury development,19,20 it also has the potential to refine or influence existing reductionist approaches that characterize traditional epidemiological inquiry. In the RRI literature, researchers have explored the attitudes and opinions regarding factors associated with running injury development. Saragiotto et al. investigated the attitudes of recreational runners on risk factors for running injury.21 These runners mainly attributed injury to training, running shoe choice, and exceeding the body´s limits. In another study among health professionals and coaches, training factors such as excessive training was found to be important risk factors.18 Finally, van Wilgen et al. investigated the attitudes among coaches and athletes from various sports.22 They found that factors related to training, situation and behavior were associated with injuries. However, only one of the coaches included in the study worked directly with runners. Based on this, there is a lack of empirical knowledge regarding the attitudes of M/L distance runners competing in national championships and their coaches on risk factors for running injury. Therefore, the purpose of the present study was to investigate the attitudes of M/L distance runners able to compete in national championships and their coaches about factors associated with running injury development. METHODS The study was designed as a survey-based study. The study is observational and therefore needs no

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permission from the system of research ethics committees according to the Danish Act on Research of Health Projects, Section 14, no. 2. Based on this, the study was conducted according to Danish law. Since no personal-identifiable information was collected, the Danish Data Protection Agency waived the request for approval, because participants indicated informed consent by clicking on a yes-button in a web-based questionnaire, and since data collections without identifiable data do not need approval in Denmark. Recruitment and data collection Data was collected in November and December 2015. The study population was M/L distance runners and their coaches who were members of a running club registered in the Danish Athletic Federation. This enabled the runners to compete in National Championships (road races, cross-country and track) and coaches to work with runners competing in National Championships. E-mails were distributed to 25 of 69 Danish Athletic Clubs registered by the Danish Athletic Federation (http://www.atletikogmotion.dk/). The 25 clubs were chosen since these clubs, according to athletic-specific homepages (http://www.statletik. dk/ and http://daf.sportstiming.dk/), had members participating in National Championships (e.g. 10 kilometer road running, 800 to 10.000 meter track, and the marathon distance) and/or local athletic events (e.g. “Bane turneringen” and Aarhus Nordic Challenge) in the prior year. It is the authors´ best guess that there were approximately 500 runners affiliated with the 25 included running clubs. The exact number was unknown given that participant recruitment occurred through contact with staff members and running coaches. This was because the Danish Athletics Federation does not have a record of its members, and so it was not possible to extract information from a known database. In the e-mail, information about the study was provided and the coaches were encouraged to respond to a coach-specific survey, whilst the runners were encouraged to respond to an athlete-specific survey through a link to a web-based survey. In addition to the e-mails, information about the study and links to surveys were posted via clubspecific social media groups (e.g. Facebook). In the event that no responses were submitted one week after the first contact, a reminder-e-mail was sent.

After an additional week of no response, the chairman / chairwomen from the club was contacted by telephone and encouraged to inform the members about the study. Runners eligible for inclusion were: (i) older than 17 and younger than 51 years of age; and, (ii) athletes training at least 40 km per week. Coaches eligible for inclusion were: (i) Above the age of 18; and, (ii) an active coach for one or more M/L distance runners competing at an athletic level. Runners or coaches under the age of 18 were excluded since parental consent was needed in case runners and coaches below 18 years were to participate. Survey Both surveys were written in Danish. In October 2015, these surveys were pilot tested amongst four runners and three coaches from one athletic club. The final surveys contained 26 questions including: (i) demographic information (age, gender, weight, height); (ii) training characteristics (e.g. average running distance, number of high-intensive sessions, running frequency per week years as active athletic runner); (iii) injury status (currently injured, current- and previous injury, previously injured, never had an injury); and, (iv) attitudes regarding injury occurrence. A RRI was defined as “a prolonged pain, which leads the runners to reduce training or competition for at least three weeks”. The main question addressed by the survey was: “Which factors do you believe influence running injury risk?”. For this question, the coach or athlete was presented with a list of 19 items (Table 1). For each item, they were asked to answer, by clicking a yes or no button, if they believed that a given factor influenced RRI risk. These 19 items (divided in to four categories: personal factors, behavior, shoes, and training-related factors) were chosen based on the findings from recent (at the time) systematic reviews 23,24 and attitudes amongst recreational runners 21. In addition, the responders had the possibility to click “other” and add additional items. Secondary questions were presented to investigate which types of RRIs the runners and coaches felt were most common. Furthermore, to increase the insight on attitudes regarding advice which could lead to running injury prevention, the coaches and athletes were asked in an open-ended question:

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Table 1. Possible answers to: “Which factors do you believe inﬂuence running injury risk?” In this question, the coach or runner was presented with a list of 19 items. For each item, they were asked to answer by clicking yes or no on the survey, if they believed that particular factor inﬂuenced running injury risk. In addition, the respondents were able to add other reasons.

“How would you advise a runner to act in order to prevent injuries?” Statistics Continuous data are presented as mean and standard deviation (SD) if they were normally distributed and as median and inter-quartile range if they did not follow a normal distribution. Categorical data are presented as numbers (n) and proportion (%). Differences in the opinions between coaches and athletes were analyzed using students’ t-test for continuous data and as chi-square test for categorical variables. Differences were considered statistically significant at p<0.05. RESULTS A total of 83 athletes (45 males, 37 females) representing 11 of the 25 invited clubs (44%) completed the survey. However, 15 (18%) of these were excluded, as they did not meet the inclusion criteria, due to not meeting minimum age requirements (n=4, 27%), low training volume (n=6, 40%), or missing several required questions and misreported answers (n=5,

33%). Eventually, 68 athletes were included in the study (Table 2). The demographic characteristics of the 19 coaches representing 11 clubs (44%) are presented in Table 3. The attitudes amongst M/L runners and their coaches, as well as factors on which they agree and disagree in relation to injury risk, are presented in Table 4. According to both M/L runners and coaches, the injury types believed to be most common were medial tibial stress syndrome, Achilles tendinopathy and iliotibial band syndrome. Regarding the open-ended question “How would you advise a runner to act in order to prevent injuries?”, the most commonly advised intervention was strength training as advised by nine athletes (13%) and two coaches (11%). Alternative training (e.g. deep-water running, cross training, swimming) was advised by seven athletes (10%). Next, eight athletes (11%) and two coaches (11%) advised avoiding increasing the intensity and/or volume too fast, and finally three athletes (4%) and two coaches (11%) emphasized the importance of communication between the coach, physiotherapist, and athlete.

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Table 2. Injury history, demographic- training-related characteristics of the included M/L athletes. Values represent mean ±SD or number and proportion. a = data presented as mean and standard deviation. b = data reported as median and interquartile range since data was not normally distributed.

Table 3. Demographic- and coaching experience of the included coaches. Values represent mean ±SD or number and proportion

DISCUSSION This is the first study to examine the attitudes of M/L distance runners able to compete in National Championships and their coaches regarding their attitudes regarding risk factors for RRI. Runners reported ignoring pain, insufficient recovery between running sessions, lack of strength, training experience and previous injury to be main risk factors, while the coaches reported ignoring pain, lack of strength, high training volume, insufficient recovery between running sessions, and stress to be the most important risk factors for running injury. According to most runners and coaches, ignoring pain stands out as the main risk factor associated with injury development. This is interesting since Jacobsson et al., in a study of Swedish elite athletes, found half of the injured athletes experienced pain one to two weeks before injury was reported.2 Attention towards early symptoms and pain management might be crucial in running injury prevention for this population of runners. In particular, amongst those

with a “no pain, no gain” attitude, since Saragiotto et al. found coaches and health professionals to report increased risk of injury development amongst such types of athletes.18 A total of 79% of the coaches and 44% of the runners reported reduced strength as another risk factor associated with increased injury risk. This finding is in accordance with the attitudes among recreational runners who also reported “lack of strength” as a risk factor for injury.21 A recently published systematic review on risk factors for RRI found only few published studies examining the role of strength-based resistance training on injury development.10 Fortunately, more scientific focus on the role of strength training on injury development is occurring.25 In epidemiological studies,10 conflicting results exist regarding the role of running frequency and risk of RRI. In the present study, insufficient recovery between running sessions was highlighted as impor-

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Table 4. Attitudes among middle- and long (M/L) distance runners and their coaches on risk factors for running-related injury. P-value is based on a test for similar proportions between runners and coaches. CI = Confidence interval. Shoe and foot posture = Running in running shoes, which do not fit foot posture (neutral, pronation). Fast progression = Fast progression in distance and/or intensity. Running style = striking at fore-foot, mid-foot and/or rear-foot. Insufficient recovery = short time between running sessions.

tant by 59% of the runners and 68% of the coaches. Still, there is little to no scientific agreement regarding the number of running sessions acceptable for different runners of different shapes and sizes. Most likely, runners with a high experience who have been accustomed to many weekly running sessions might not be as prone to injury compared with low experience runners running a few sessions per week and then suddenly changing to more sessions per week.26 Running “too much, too soon” has been discussed in the scientific literature,11,23,27 since overuse-related

injuries, theoretically, are a result of training errors.28 Interestingly, the great interest amongst scientists on the deleterious role of excessive running29,30 does not seem to reflect the attitudes amongst the runners, since below 15% found excessive training distance and/or intensity to be risk factors for injury. Conversely, 74% of the coaches found excessive distance as being associated with injury risk. Clearly, this shows a gap in attitudes between coaches on one side and the athletes on the other. Interestingly, only two coaches advised avoiding increasing the intensity and/or volume too fast in order to prevent injuries. This could lead to the assumption, that

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coaches are unfamiliar with, or do not prioritize, the proposed link between risk factors for injury and targets for intervention strategies. If the prevention is better made by knowledge about risk factors,9 more coaches should pay attention to excessive training avoidance and pain management as preventive strategies. Unfortunately, the evidence-base for determining the appropriate dose of running for runners with different characteristics is non-existent. More studies are needed to identify training schedules associated with a low injury risk for different types of runners. The study has its limitations since the survey could have been more nuanced providing respondents with the possibility to address each question using a Likert scale ranging from “no importance” to “high importance” rather than the dichotomized approach used in the present study. Unfortunately, it was not possible to create a Likert scale in the web-based system used to set-up the survey. Therefore, a dichotomized solution was used. In addition, the sample size may be a limitation since the beliefs of the 68 runners included might not reflect the belief of the approximately 500 runners (14%) who were members of the 25 clubs, which were contacted. Consequently, the proportions reported in this article may be over- and/or underestimated because of selection problems. In addition, the choice of injury definition in the present study might be considered as a limitation, since usage of the consensus-based definition of running injury proposed by Yamato et al.31 could reveal other results than those presented. CONCLUSIONS The primary purpose of the present study was to investigate the attitudes of M/L distance runners and their coaches about factors associated with RRI development. M/L distance runners and their coaches report “ignoring pain” as a major risk factor for running injury development, while a majority of the coaches reported “reduced muscle strength” and “high running distance” to be associated with injury. The need to further investigate the athlete’s psychological profile in relation to runningrelated injury development is warranted based on the novel results generated in this study. In and of itself, ‘ignoring pain’ is not particularly informative when thinking about injury prevention solutions.

However, it is more than likely a proxy indication for another exposure, since behavior itself is unable to cause running-related injury. An athlete needs to run to sustain an injury. Therefore, the proxy variable might be the amount of running participation that might be readily quantifiable and meaningful in practical terms. The questions remains how much running, if any at all, runners having slight symptoms or mild pain, are able to tolerate before these symptoms develop into a RRI. REFERENCES 1. Timpka T, Alonso JM, Jacobsson J, et al. Injury and illness definitions and data collection procedures for use in epidemiological studies in Athletics (track and field): consensus statement. Br J Sports Med. 2014; 48(7):483-490. 2. Jacobsson J, Timpka T, Kowalski J, et al. Injury patterns in Swedish elite athletics: annual incidence, injury types and risk factors. Br J Sports Med. 2013; 47(15):941-952. 3. Jacobsson J, Timpka T, Kowalski J, et al. Prevalence of musculoskeletal injuries in Swedish elite track and field athletes. Am J Sports Med. 2012; 40(1):163-169. 4. Huxley DJ, O’Connor D, Healey PA. An examination of the training profiles and injuries in elite youth track and field athletes. Eur J Sport Sci. 2014; 14(2):185-192. 5. Nielsen RO, Ronnow L, Rasmussen S, et al. A prospective study on time to recovery in 254 injured novice runners. PLoS One. 2014; 9(6):e99877. 6. D’Souza D. Track and field athletics injuries--a one-year survey. Br J Sports Med. 1994; 28(3):197-202. 7. Bennell KL, Malcolm SA, Thomas SA, et al. Risk factors for stress fractures in track and field athletes. A twelve-month prospective study. Am J Sports Med. 1996; 24(6):810-818. 8. Yeung SS, Yeung EW, Gillespie LD. Interventions for preventing lower limb soft-tissue running injuries. Cochrane Database Syst Rev 2011; (7):CD001256. doi(7):CD001256. 9. Finch C. A new framework for research leading to sports injury prevention. J Sci Med Sport. 2006; 9(1-2):3-9. 10. Hulme A, Nielsen RO, Timpka T, et al. Risk and protective factors for middle- and long-distance running-related injury: A systematic review. Sports Med. 2016;:In press. 11. Malisoux L, Nielsen RO, Urhausen A, et al. A step towards understanding the mechanisms of runningrelated injuries. J Sci Med Sport. 2015; 18(5):523-528.

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12. Hulme A, Finch CF. From monocausality to systems thinking: a complementary and alternative conceptual approach for better understanding the development and prevention of sports injury. Inj Epidemiol. 2015; 2(1):31. 13. Herzog W. The Problem with Running Injuries. J Sport Health Sci. 2016; 5(2):171. 14. Hulme A, Finch C. The epistemic basis of distance running injury research: A historical perspective. J Sport Health Sci. 2016; 5(2):172-175. 15. Morgan DL. Paradigms lost and pragmatism regained. Methodological implications of combining qualitative and quantitative methods. J Mixed Meth Res. 2009; 1(1):48-76. 16. van Griensven H., Freshwater D. The 1 + 1 = 3 integration challenge. J Mixed Meth Res. 2014; 19(5):367-371. 17. Maxwell JA. The importance of qualitative research for causal explanation in education. Qualitative Inquiry. 2012; 8(8):655-661. 18. Saragiotto BT, Di Pierro C, Lopes AD. Risk factors and injury prevention in elite athletes: a descriptive study of the opinions of physical therapists, doctors and trainers. Braz J Phys Ther. 2014; 18(2):137-143. 19. Johnson U. Athletes’ experiences of psychosocial risk factors preceding injury. Qual Res Sport, Exerc Health. 2011; 3(1):99-115. 20. Johnson U. Psychosocial antecedents of sport injury, prevention, and intervention: An overview of theoretical approaches and empirical ﬁndings. Int J Sport Exerc Psych. 2007; 5(4):352-369. 21. Saragiotto BT, Yamato TP, Lopes AD. What do recreational runners think about risk factors for running injuries? A descriptive study of their beliefs and opinions. J Orthop Sports Phys Ther. 2014; 44(10):733-738. 22. van Wilgen CP, Verhagen EA. A qualitative study on overuse injuries: the beliefs of athletes and coaches. J Sci Med Sport. 2012; 15(2):116-121.

23. Nielsen RO, Buist I, Sorensen H, et al. Training errors and running related injuries: a systematic review. Int J Sports Phys Ther. 2012; 7(1):58-75. 24. Saragiotto BT, Yamato TP, Hespanhol Junior LC, et al. What are the main risk factors for running-related injuries? Sports Med. 2014; 44(8):1153-1163. 25. Baltich J, Emery CA, Whittaker JL, et al. Running injuries in novice runners enrolled in different training interventions: a pilot randomized controlled trial. Scand J Med Sci Sports. 2016: In press. 26. Nielsen RO, Malisoux L, Møller M, et al. Shedding light on the etiology of sports injuries: A look behind the scenes of time-to-event analyses. J Orthop Sports Phys Ther. 2016; 46(4):300-311. 27. Soligard T, Schwellnus M, Alonso JM, et al. How much is too much? (Part 1) International Olympic Committee consensus statement on load in sport and risk of injury. Br J Sports Med. 2016; 50(17):10301041. 28. Hreljac A. Etiology, prevention, and early intervention of overuse injuries in runners: a biomechanical perspective. Phys Med Rehabil Clin N Am. 2005; 16(3):651-67. 29. Nielsen RO, Parner ET, Nohr EA, et al. Excessive progression in weekly running distance and risk of running-related injuries: an association which varies according to type of injury. J Orthop Sports Phys Ther. 2014; 44(10):739-747. 30. Nielsen RO, Cederholm P, Buist I, et al. Can GPS Be Used to Detect Deleterious Progression in Training Volume Among Runners? J Strength Cond Res 2013; 27(6):1471-1478. 31. Yamato TP, Saragiotto BT, Lopes AD. A consensus deﬁnition of running-related injury in recreational runners: a modiﬁed Delphi approach. J Orthop Sports Phys Ther. 2015; 45(5):375-380.

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IJSPT

ORIGINAL RESEARCH

CAVITATION SOUNDS DURING CERVICOTHORACIC SPINAL MANIPULATION James Dunning, DPT, MSc, FAAOMPT1,2 Firas Mourad, PT, OMPT, Dip. Osteopractic1,2 Andrea Zingoni, MSc3 Raffaele Iorio, PT, Cert. DN4 Thomas Perreault, DPT, OCS, Dip. Osteopractic2 Noah Zacharko, DPT, FAAOMPT, Dip. Osteopractic2,5 César Fernández de las Peñas, PhD, DO, PT6 Raymond Butts, PhD, DPT, MSc2,7 Joshua A. Cleland, PhD, DPT, FAAOMPT8

ABSTRACT Background: No study has previously investigated the side, duration or number of audible cavitation sounds during high-velocity low-amplitude (HVLA) thrust manipulation to the cervicothoracic spine. Purpose: The primary purpose was to determine which side of the spine cavitates during cervicothoracic junction (CTJ) HVLA thrust manipulation. Secondary aims were to calculate the average number of cavitations, the duration of cervicothoracic thrust manipulation, and the duration of a single cavitation. Study Design: Quasi-experimental study Methods: Thirty-two patients with upper trapezius myalgia received two cervicothoracic HVLA thrust manipulations targeting the right and left T1-2 articulation, respectively. Two high sampling rate accelerometers were secured bilaterally 25 mm lateral to midline of the T1-2 interspace. For each manipulation, two audio signals were extracted using Short-Time Fourier Transformation (STFT) and singularly processed via spectrogram calculation in order to evaluate the frequency content and number of instantaneous energy bursts of both signals over time for each side of the CTJ. Result: Unilateral cavitation sounds were detected in 53 (91.4%) of 58 cervicothoracic HVLA thrust manipulations and bilateral cavitation sounds were detected in just five (8.6%) of the 58 thrust manipulations; that is, cavitation was significantly (p<0.001) more likely to occur unilaterally than bilaterally. In addition, cavitation was significantly (p<0.0001) more likely to occur on the side contralateral to the clinician’s short-lever applicator. The mean number of audible cavitations per manipulation was 4.35 (95% CI 2.88, 5.76). The mean duration of a single manipulation was 60.77 ms (95% CI 28.25, 97.42) and the mean duration of a single audible cavitation was 4.13 ms (95% CI 0.82, 7.46). In addition to single-peak and multipeak energy bursts, spectrogram analysis also demonstrated high frequency sounds, low frequency sounds, and sounds of multiple frequencies for all 58 manipulations. Discussion: Cavitation was significantly more likely to occur unilaterally, and on the side contralateral to the short-lever applicator contact, during cervicothoracic HVLA thrust manipulation. Clinicians should expect multiple cavitation sounds when performing HVLA thrust manipulation to the CTJ. Due to the presence of multi-peak energy bursts and sounds of multiple frequencies, the cavitation hypothesis (i.e. intra-articular gas bubble collapse) alone appears unable to explain all of the audible sounds during HVLA thrust manipulation, and the possibility remains that several phenomena may be occurring simultaneously. Level of Evidence: 2b Key words: Cavitation, cervicothoracic spine, spinal manipulation, thrust

PhD student, Escuela Internacional de Doctorado, Universidad Rey Juan Carlos, Alcorcón, Madrid, Spain 2 American Academy of Manipulative Therapy Fellowship in Orthopaedic Manual Physical Therapy, Montgomery, AL, USA 3 Università di Pisa, Information Engineering Department, Pisa, Italy 4 Koinè Physiotherapy, Florence, Italy 5 Osteopractic Physical Therapy of the Carolinas, Charlotte, NC, USA 6 Department of Physical Therapy, Occupational Therapy, Rehabilitation and Physical Medicine, Universidad Rey Juan Carlos, Alcorcón, Madrid, Spain 7 Research Physical Therapy Specialists, Columbia, SC, USA 8 Franklin Pierce University, Manchester, NH, USA

Acknowledgements None of the authors received any funding for this study. The authors wish to thank all the participants of the study. Declaration of Interests The authors report no declarations of interest.

CORRESPONDING AUTHOR Dr. James Dunning 1036 Old Breckenridge Lane Montgomery, AL 36117 E–mail: jamesdunning@hotmail.com

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INTRODUCTION Reductions in pain and disability following highvelocity low-amplitude (HVLA) thrust manipulation to the cervicothoracic region have been widely reported in patients with neck pain1-11 and shoulder pain.12-16 However, the frequency, location, and etiology of the cracking, popping or clicking noises that often accompany HVLA thrust manipulative procedures to the spine17-25 are still poorly understood.23,26-30 Four previous studies31-34 have suggested that the “audible popping” following HVLA thrust manipulation is not related to the clinical outcomes of pain and/or disability. Nevertheless, many clinicians19,22,35 and researchers20,21,36-42 still appear to repeat the HVLA thrust manipulation if they do not hear or palpate popping sounds. Moreover, Evans and Lucas27 proposed the “audible popping”, or the “mechanical response” that “occurs within the recipient”, should be present to satisfy the criteria for a valid manipulation.27 However, it remains to be elucidated whether HVLA thrust manipulation to the cervicothoracic spine should normally be accompanied by single, multiple or no cavitation sounds. Furthermore, understanding whether the cavitation phenomenon during cervicothoracic HVLA thrust manipulation is an ipsilateral, contralateral or bilateral event may help inform clinicians in selecting the appropriate manipulation technique that will most effectively target the dysfunctional articulation with the ultimate goal of reducing pain and disability. The traditional expectation of a single pop or cavitation sound emanating from the target or dysfunctional facet joint during HVLA thrust manipulation43,44 is not consistent with the existing literature for the upper cervical,26 lower cervical,24,45 thoracic25 or lumbar17,20,25 regions. Moreover, the evidence suggests that HVLA thrust manipulation directed at the spine creates multiple cavitation sounds.17,24-26,45 Nevertheless, the question of whether these multiple cavitation sounds emanate from the same joint, adjacent ipsilateral or contralateral joints, or even extra-articular soft-tissues remains to be elucidated.17,18,20,25,26,46 To date, only three studies18,24,26 have investigated the side of joint cavitation during cervical spine manipulation. During “lateral to medial and rotatory” HVLA thrust manipulations targeting the C3-4 facet joint, Reggars and Pollard24 found 47 (94%) of

50 subjects exhibited “cracking sounds” on the contralateral side to the applicator contact, while two subjects exhibited bilateral sounds and one subject an ipsilateral sound. Additionally, following C3-4 thrust manipulations in 20 asymptomatic subjects, Bolton et al18 reported cavitation sounds were significantly more likely to occur on the contralateral side to the applicator for “rotation” manipulations, but equally likely to occur on either side during “sidebending” manipulations. Nevertheless, Bolton et al18 made the assumption that the side with the larger amplitude sound wave was the side of “initial cavitation” and hence did not report if single or multiple cavitations occurred. That is, unless single cavitation events occurred during all cervical manipulations, which is unlikely given the findings of previous studies,17,20,24,25,45 the possibility remains that the “initial cavitation” occurred on one side, and additional cavitations that were not counted also occurred ipsilaterally and/or contralaterally. Most recently, Dunning et al26 reported bilateral cavitation sounds in 34 (91.9%) of 37 manipulations, while unilateral cavitation sounds were detected in just 3 (8.1%) manipulations following HVLA thrust manipulation targeting the upper cervical spine (C1-2) articulation. However, it is unknown if the same findings would occur in a different spinal region—i.e. the cervicothoracic junction (CTJ)—and whether using a different HVLA thrust technique with the patient in prone, that is traditionally considered a “lateral break” manipulation35,47-49 (due to the simultaneous delivery of lateral flexion and lateral translation forces as opposed to primarily employing rotatory forces for the thrusting impulse35,50), would alter the side of cavitation and therefore the location of the target articulation that will most likely be effected by the high-velocity thrusting forces.17,18,20,24,26,28 Gas bubble collapse,51 or the cavitation phenomenon, has been traditionally accepted as the mechanism for creating the joint cracking sound.18,23,27,30,45,51-53 However, a recent study by Kawchuk et al29 challenged the cavitation hypothesis, and proposed that joint cracking is associated with cavity formation within synovial fluid rather than cavity collapse. Nevertheless, although this first in-vivo macroscopic demonstration of tribonucleation was recorded using rapid cine magnetic resonance images on 10 MCP joints, it was from a single subject.29 Furthermore, the notion

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that the audible popping sounds were coming from cavity inception, rather than collapse of a pre-existing bubble, is not new and was first proposed by Roston and Haines as early as 1947.54 However, neither of these two studies29,54 can be generalized to zygapophyseal joints. Identifying normative values for the duration of HVLA thrust procedures23,26,55,56 for different spinal regions may help facilitate a better understanding of the physical parameters surrounding spinal manipulation28,50,55 (e.g. velocity, acceleration) and the specific psychomotor skills required by practitioners to efficiently perform spinal thrust manipulations.28,48 Additionally, it still remains to be elucidated whether the popping sounds during HVLA thrust manipulation originate from intra-articular gas bubble collapse, cavity inception within synovial fluid, or extra-articular events.26,29,30,46,51 Therefore, identifying the duration of individual cavitation sounds,23,24,26 analyzing the instantaneous energy bursts and frequency content of the sound waves23,26,45 produced during thrust manipulations may help uncover the etiology29,46,52,53,57—i.e. what structures, tissues, or mechanisms are involved—and therefore the relative importance of the audible sounds during thrust manipulations.27,31,32,34 For cervical manipulations, the duration of the thrusting procedure has been reported to be 80-200 ms.23,26,50,55,56 Additionally, using 95% of the instantaneous energy burst—i.e. the amount of energy released in a given sampling interval of the spectrogram—to calculate the duration of single cavitation sounds during upper cervical HVLA thrust manipulation, Dunning et al26 reported a mean duration of 5.66 ms. However, no study has previously measured the duration of the thrusting procedure or the duration of single cavitation sounds, during HVLA thrust manipulation to the CTJ. To the best of the authors’ knowledge, no study has investigated the side, duration or number of audible cavitation sounds during HVLA thrust manipulation to the cervicothoracic spine. Therefore, the primary purpose of the study was to determine which side of the spine cavitates during cervicothoracic HVLA thrust manipulation. Secondary aims of the study were to calculate the duration of a single cervicothoracic thrust manipulation procedure, and the

average number of cavitation sounds following cervicothoracic HVLA thrust manipulation. METHODS Participants Thirty-two individuals with upper trapezius myalgia, i.e. a painful upper trapezius muscle, (20 females and 12 males) were recruited by convenience sampling from a private physical therapy outpatient clinic in Florence, Italy during November of 2013. Their ages ranged between 23 and 65 years with a mean (SD) of 39 (11) years. Height ranged between 152 and 182 cm with a mean (SD) of 170.1 (8.5) cm. Weight was 50.0 kg to 96.0 kg with a mean (SD) of 67.7 (12.6) kg. All subjects reported being physically active, to include walking, running, cycling or regular sports participation. For subjects to be eligible, they had to present with neck pain for greater than three months, have a primary complaint of a painful spot (i.e., active trigger point) in the upper trapezius muscle, and be between 18 and 65 years of age. The ethics committee at the Universidad Rey Juan Carlos, Madrid, Spain, approved this study. All subjects provided written informed consent before their participation in the study. Patients were excluded if they exhibited: 1) any red flags (i.e., tumor, fracture, metabolic diseases, rheumatoid arthritis, osteoporosis, resting blood pressure greater than 140/90 mmHg, prolonged history of steroid use, etc.); 2) presented with 2 or more positive neurologic signs consistent with nerve root compression (muscle weakness involving a major muscle group of the upper extremity, diminished upper extremity deep tendon reflex, or diminished or absent sensation to pinprick in any upper extremity dermatome); 3) presented with a diagnosis of cervical spinal stenosis; 4) exhibited bilateral upper extremity symptoms; 5) had evidence of central nervous system disease (hyperreflexia, sensory disturbances in the hand, intrinsic muscle wasting of the hands, unsteadiness during walking, nystagmus, loss of visual acuity, impaired sensation of the face, altered taste, the presence of pathological reflexes); 6) had a history of whiplash injury within the previous three months; or, 7) had prior surgery to the neck or thoracic spine. Of the 33 patients that were

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invited to enter the study, none refused participation; however, one subject was excluded due to a history of a previous whiplash injury. Notably, pain or disability scores were not collected in any subjects for two reasons: (1) the primary purpose of this study was to investigate the frequency, location and possible etiologies of the cavitation phenomenon during cervicothoracic HVLA thrust manipulation at a single point in time (i.e. no follow-up period), not to measure changes in pain or disability over time in response to a single manipulation technique given on just one occasion, and (2) all subjects were current patients at a physiotherapy practice in Florence, Italy, and as such, were already receiving conventional physiotherapy treatments for their primary complaint of upper trapezius myalgia. Moreover, significant reductions in pain and disability scores following HVLA thrust manipulation to the cervicothoracic region have already been widely investigated and reported in patients with neck pain1-11 and shoulder pain.12-16 However, although cracking, popping or clicking noises often accompany HVLA thrust manipulative procedures,17-25 the frequency, location and etiology of the cavitation phenomenon itself is still poorly understood.23,26-30 Manipulative Physiotherapist A single, U.S. licensed physical therapist performed all of the cervicothoracic HVLA thrust manipulations in the current study. At the time of data collection, the physical therapist had completed a post-graduate Master of Science in Advanced Manipulative Therapy, had worked in clinical practice for 14 years, and routinely used cervicothoracic HVLA thrust manipulation in daily practice. Cervicothoracic Junction (CTJ) HVLA Thrust Manipulation Technique A single “lateral break” HVLA thrust manipulation directed to the CTJ with the patient prone was performed (Figure 1). T1-2 was the target level because this segment is in the center of the three articulations (i.e. C7-T1, T1-T2, T2-3) that are considered to be primarily affected by the manual forces during prone HVLA thrust manipulations to the CTJ.12,22,28,47,50,58,59 For this technique,47 the short or lower lever was produced by having the therapist’s proximal phalanx, metacarpal, web space and

Figure 1. High-velocity low-amplitude thrust manipulation directed to the articulation of the left cervicothoracic (T1-2) junction.

thumb of the right hand contact the superomedial aspect of the patient’s right shoulder girdle. The long or upper lever was manufactured by having the therapist place the heel and palm of his left hand over the temporal region of the patient’s lateral cranium. To localize the forces to the left T1-2 articulation, secondary levers of extension, lateral flexion, translation and minimal rotation were used. While maintaining the secondary levers, the therapist performed a single HVLA thrust manipulation using the simultaneous delivery of the thrusting primary levers of lateral flexion from the upper lever and lateral translation from the lower lever, i.e., a lateral break. This was repeated using the same procedure but directed to the right T1-2 articulation. Prior to data collection, an independent researcher made random allocation cards using a computer-generated table of randomly assigned numbers;60 these cards were then used to determine the target side and delivery order of the T1-2 HVLA thrust manipulations for all subjects. Cavitation sounds—i.e. popping or cracking noises—were heard on all HVLA thrust manipulations; hence, there was no need for second attempts.

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Accelerometer Placement and Sound Collection Prior to the delivery of cervicothoracic HVLA thrust manipulation, skin mounted accelerometers were secured bilaterally 25 mm lateral to the midline of the T1-T2 interspace (Figure 2). The microphones were connected to a data acquisition system (FOCUSRITE, High Wycombe, Buckinghamshire, U.K., Scarlett 2i2, 96 KHz, 24-bit conversion) and a MacBook Pro laptop with AUDACITY software (Open Source Software, Carnegie Mellon University, Pennsylvania, U.S.A.) for audio acquisition.26 Sampling frequency was set at 96,000 Hz and the amplitude was normalized by AUDACITY software to values ranging between -1 and +1 (no unit of measurement). With the order of delivery randomized (i.e. right side versus left side), all subjects then received two HVLA thrust manipulations: one targeting the left (T1-2) CTJ, and one targeting the right (T1-2) CTJ. The sound wave signals and resultant cavitation sounds during the cervicothoracic HVLA thrust manipulations were recorded by an individual not involved in data extraction or analysis. Data extraction and processing occurred later and were performed by an individual blinded to target side. Although target side and delivery order were randomly assigned using a computergenerated table of randomly assigned numbers, it was not possible to fully blind the third researcher who performed data analysis because knowledge of target side was required to complete some of the statistical testsâ&#x20AC;&#x201D;for example, whether cavitation was more likely to occur on the side ipsilateral or contralateral to the clinicianâ&#x20AC;&#x2122;s short-lever applicator. Data Extraction Short-Time Fourier Transformation (STFT) was used to process the sound signals and obtain spectrograms

Figure 2. Bilateral placement and securing of skin-mounted accelerometers 25 mm lateral to the midline of the T1-2 interspace.

for each thrust manipulation.26 A spectrogram is a two-dimensional representation of a signal with time on the x-axis, frequency on the y-axis, and color as a third dimension to express the amplitude, or power of the sound (Figure 3). For each two-channel audio recording, the spectrograms were computed using STFT in order to evaluate the frequency content of both signals over time. The epoch length was set to 0.78 ms (i.e. 75 times the sampling rate) with a 0.1% overlap between adjacent epochs, resulting in a frequency resolution of 94 Hz. The frequency scale was set between 10 Hz and 23 kHz, since this is the audible spectrum for a human being (including a small margin of error).61 Data Processing The sound in every audio track was processed as a digital signal with the amplitude varying discretely as a function of time. Each channel was depicted by a separate graph, representing the two recordings of the left and right accelerometers during a single HVLA thrust manipulation to the CTJ. Although the recordings were collected and processed singularly for each person and for each manipulation, we did jointly inspect and analyze the left and right channels for each HVLA thrust manipulation in order to determine whether the cavitation phenomenon was a bilateral, ipsilateral or contralateral event, and in order to accurately sum the total number of cavitations (i.e. pops) during a single manipulation. In order to isolate the time interval in which the manipulation took place, the audio tracks of the left and right channels (relative to a single manipulation) were first listened to using a stereophonic system. The peculiar sound emitted, together with visual inspection of the right and left graphs of the digital audio signal, allowed for easy recognition of such an interval. The correct time interval featuring the manipulation event was then confirmed and adjusted by decelerating the audio speed by a factor of 0.01 and listening to the track again. This allowed us to identify the beginning and the end of the manipulations (based on sound, not angular movements of the spine), and also to identify how many cavitations (i.e. pops) were present. More specifically, this operation permitted us to increase the resolution of the human ear by 100 fold, allowing us to discriminate and sum the total number of cavita-

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tions. Moreover, listening of the audio tracks with a 100-fold deceleration factor and visual inspection of the spectrograms for peaks were both used to determine the number of cavitations present. The spectrograms show the â&#x20AC;&#x153;locationâ&#x20AC;? of the energy of the audio signals over time and over frequency jointly. In figure 3, the spectrograms for the right and left channels for a single HVLA thrust manipulation are depicted, with time on the x-axis, frequency on the y-axis and energy on the z-axis (using a map of colors). Process for Counting the Number of Cavitations Per the protocol previously described,26 the graphs representing the amount of released energy over time in both the left and right accelerometry chan-

nels were visually inspected in order to identify instantaneous energy bursts corresponding to cavitations (Figure 4). The total number of cavitations per manipulation was the sum of the number of energy bursts identified. Process for Determining the Side of Cavitation For each of the 252 pops generated during 58 cervicothoracic HVLA thrust manipulations, the side of cavitation was determined by inspecting each of the energy bursts for the right and left spectrograms.26 Since graphs were computed and the amount of energy was quantified at each epoch separately for the two channels, the side of cavitation could be immediately determined by looking at which side the energy burst occurred on. In the event of

Figure 3. Spectrograms for the left and right audio channels during cerviothoracic HVLA thrust manipulation. Vertical energy peaks represent individual pops.

Figure 4. Amount of energy released over time for the right and left accelerometry channels. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 647

simultaneous bursts on both channels, the one that began earlier and had the higher energy value was selected; moreover, the smaller and delayed energy burst represented the echo of the original event. A similar methodology for determining the side of cavitation was previously reported for the upper cervical spine.26 Process for Calculating the Duration of a Single Cavitation For each of the 252 cavitations (i.e. popping sounds) detected during 58 cervicothoracic HVLA thrust manipulations, the time interval between the beginning of the ascent of the first energy burst and the end of the descent of the last energy burst of a cavi-

tation event was considered as the duration of a single cavitation (Figure 5). Process for Calculating the Duration of the Thrust Manipulation For each thrust manipulation, the time interval between the beginning of first cavitation and the end of the last cavitation was considered as the duration of the thrusting procedure (Figure 6).26 However, we did not measure the actual forces against time; therefore, the duration of the thrust manipulation likely does not include the time from when the force beyond the preload first began to be applied, or the entire interval from when the peak forces dropped back to zero.28,48,50

Figure 5. The time interval used to calculate the duration of a single pop during cervicothoracic HVLA thrust manipulation.

Figure 6. The time interval used to calculate the duration of cervicothoracic HVLA thrust manipulation. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 648

Data Analysis Sound waves resulting from the cervicothoracic HVLA thrust manipulations were displayed in graphical format. Each subject had one right and one left graph corresponding with each thrust procedure (i.e. four graphs in total for each subject). Means and standard deviations were calculated to summarize the average number of pops, the duration of cervicothoracic thrust manipulation, and the duration of a single cavitation. The primary aim, to determine which side of the spine cavitates during CTJ (T12) HVLA thrust manipulation, was examined using a Chi-square test. The probability for unilateral or bilateral cavitation events was calculated using the binomial test assuming an expected probability of 50% (i.e. a reference proportion of 0.5). Data analysis was performed using SPSS 23.0. RESULTS Subjects ranged between 23 and 65 years of age, with a mean of 39 (SD: 11) years. Of the 252 total cavitations during 58 HVLA thrust manipulations, 22 occurred ipsilateral and 230 occurred contralateral to the targeted T1-2 articulation; that is, cavitation was significantly more likely to occur on the side contralateral to the short-lever applicator of the manipulative physiotherapist (p<0.0001) following right or left thrust manipulation to the CTJ. Moreover, during T1-2 HVLA thrust manipulation targeting the right or left CTJ, the resulting cavitation sounds were 10.5 times more likely to occur on the side contralateral to the short-lever applicator of the manipulative physiotherapist than the ipsilateral side. All 58 cervicothoracic HVLA thrust manipulations resulted in one or more audible joint cavitation sounds (range, 1-8). Two hundred fifty-two cavitation sounds were detected following 58 cervicothoracic thrust manipulations giving a mean of 4.35 (95%CI 2.88, 5.76) distinct cavitation sounds (i.e. pops or cracks) per cervicothoracic HVLA thrust manipulation procedure. More specifically and on average, for each cervicothoracic HVLA thrust manipulation procedure, 3.97 (SD 1.65) of the 4.35 cavitation sounds (i.e. 91.3%) occurred on the side contralateral to the short-lever applicator of the physiotherapist, whereas, 0.38 (SD 0.75) of the 4.35 cavitation sounds occurred ipsilateral (i.e. 8.7%).

Unilateral cavitation sounds were detected in 53 (91.4%) of the 58 cervicothoracic lateral break HVLA thrust manipulations and bilateral cavitation sounds were detected in just 5 (8.6%) of the 58 thrust manipulations; that is, cavitation was significantly (binomial Test, p<0.001) more likely to occur unilaterally than bilaterally. One distinct cavitation sound (i.e. a single popping noise) was produced in 4 (6.9%) of the manipulations, whereas 2 (3.5%), 12 (20.7%), 10 (17.2%), 15 (25.9%), 13 (22.4%), 1 (1.7%) and 1 (1.7%) manipulations produced 2, 3, 4, 5, 6, 7 and 8 distinct cavitation sounds, respectively. The mean duration of a single cavitation was 4.13 ms (95%CI: 0.82, 7.46) and the mean duration of a single CTJ HVLA thrust manipulation was 60.77 ms (95%CI 28.25, 97.42). In addition to single-peak and multi-peak energy bursts, high frequency sounds, low frequency sounds, and sounds of multiple frequencies for each of the 58 cervicothoracic HVLA thrust manipulations were also identified via spectrogram analysis (Figures 3, 5 and 6). DISCUSSION Side of the cavitation The results indicate that cavitation was significantly more likely to occur on the side contralateral to the short-lever applicator of the manipulative physiotherapist following right or left cervicothoracic HVLA thrust manipulation. In addition, unilateral cavitation sounds were detected in 53 (91.4%) HVLA thrust manipulations, while bilateral cavitation sounds were detected in just 5 (8.6%) cases. Resulting cavitation sounds were 10.5 times more likely to occur on the side contralateral to the short-lever applicator of the manipulative physiotherapist than the ipsilateral side. Understanding whether the cavitation phenomenon during cervicothoracic HVLA thrust manipulation is an ipsilateral, contralateral or bilateral event may help inform clinicians in selecting the appropriate thrust manipulation technique that will most effectively target the dysfunctional articulation with the ultimate goal of reducing pain and disability. Previous authors have investigated the frequency and location of audible cavitations during cervi-

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cal18,24,26 and lumbopelvic17,20 HVLA thrust manipulation; however, this study is the first to report the frequency and side of cavitation during cervicothoracic HVLA thrust manipulation. Additionally, in the current study accelerometers were mounted directly over the target articulation (i.e. 25 mm lateral to the midline of the T1-T2 interspace), whereas both Bolton et al18 and Reggars and Pollard24 mounted microphones over the articular pillar and transverse process, respectively, of the C2 vertebra when the target was the C3-4 articulation in each of those studies. Additionally, Bolton et al18 used a significantly lower sampling frequency of 2000 Hz (compared to 96,000 Hz in our study); thus, they were only able to analyze signal amplitude in the determination for the side of the cavitation. Furthermore, Bolton et al18 made the assumption that the side with the larger amplitude sound wave was the side of “initial cavitation” and hence did not report if single or multiple cavitations occurred. Unless single cavitation events occurred during all cervical manipulations, which is unlikely given the findings of previous studies,17,20,24,25,45 the possibility remains that the “initial cavitation” occurred on one side, and additional cavitations that were not counted also occurred ipsilaterally and/or contralaterally at adjacent segments. Notably, Ross et al25 found most thoracic and lumbar HVLA thrust manipulations produced two to six audible cavitation sounds with an average error from the target joint of 3.5 cm and 5.29 cm, respectively. Additionally, Beffa and Mathews17 reported lumbar and sacroiliac HVLA thrust manipulations had low specificity and poor accuracy for the target articulation. Of the 252 total cavitations identified in this study, 22 (8.7%) occurred ipsilateral and 230 (91.3%) occurred contralateral; that is, cavitation was significantly more likely to occur on the side contralateral to the short-lever applicator of the manipulative physiotherapist. Therefore, considering the findings of previous studies17,20,25 and based on the results of this study, in order to maximize the likelihood that the target articulation is indeed manipulated, it may be appropriate to perform the T1-2 HVLA thrust manipulation with the practitioner standing on the target side of the CTJ, i.e., the short lever applicator on the side opposite the target or symptomatic articulation.

Number of cavitations per thrust Following 58 cervicothoracic thrust manipulations, 252 cavitation sounds were identified resulting in a mean of 4.35 distinct cavitations (i.e. popping or cracking noises) and a range of one to eight cavitations per T1-2 HVLA thrust manipulation. Similarly, following 37 upper cervical thrust manipulations, Dunning et al26 reported a mean of 3.57 (range of 1 to 7) cavitations per C1-2 HVLA thrust manipulation. Likewise, Reggars45 reported 123 individual “joint cracks” resulting in a mean of 2.46 cavitations and a range of 1-5 cavitations per C3-4 HVLA thrust manipulation. Similarly and in agreement with the current study, Reggars and Pollard24 reported a mean of 2.32 (range of 1 to 5) cavitations per C3-4 manipulation. Although bubble collapse, or the cavitation model51 has been widely accepted for the past four decades as the mechanism of “joint cracking”,18,23,27,30,45,51-53 a recent study by Kawchuk et al29 reported a “dark intra-articular void” during MCP distraction. Notably, this “dark intra-articular void” was associated with concurrent sound production; that is, the “joint cracking” was associated in time with cavity formation (rather than cavity collapse) within the synovial fluid, and with an average of 1.89 mm of joint surface separation. Kawchuk et al29 referred to this process as tribonucleation; that is, when sufficient distractive force overcomes the viscous attraction or adhesive forces between opposing joint surfaces, rapid separation of the articulation occurs with a resulting drop in synovial pressure, allowing dissolved gas to come out of solution to form a bubble, cavity, clear space or void within the joint. In this study sounds composed of single energy bursts (i.e. single audible popping sounds) and also sounds composed of multiple energy bursts (i.e. multiple audible popping sounds) were observed. However, whether the multiple cavitation sounds found in this study emanated from the same joint, adjacent ipsilateral or contralateral facet or uncovertebral joints, or even extra-articular soft-tissues remains to be elucidated. In addition to single and multiple energy releases, high frequency sounds, low frequency sounds, and sounds of multiple frequencies were also identified in this study. Therefore, as opposed to the cavitation hypothesis alone

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being able to explain all of the audible sounds during HVLA thrust manipulation, the possibility remains that several phenomena may be occurring simultaneously. Notably, Shekelle57 suggested HVLA thrust manipulation may affect the following pathoanatomic lesions: (1) “release of entrapped synovial folds”, (2) “disruption of intra- or peri-articular adhesions”, (3) “unbuckling of motion segments that have undergone disproportionate displacements”, and/or (4) “sudden stretching of hypertonic muscle”.57 Duration of an individual cavitation The mean duration of a single cavitation during cervicothoracic HVLA thrust manipulation was 4.13 ms (95% CI: 0.82, 7.46) in this study. This value approximates the 4 ms duration reported by Reggars and Pollard24 for the “average length of joint crack sounds” and the 5.66 ms duration reported by Dunning et al26 for the mean duration of a “single pop” during upper cervical thrust manipulation. Nevertheless, Herzog et al23 reported triphasic “cavitation signals” with a mean duration of 20 ms, however, it is unclear whether this value represents single or multiple cavitation sounds. Unlike previous studies,23,24,26 the time interval between the beginning of the ascent of the first energy burst and the end of the descent of the last energy burst of a cavitation event was calculated and used for the duration of a single pop in this study. Therefore, the interval was representative of the duration of 252 individual cavitation sounds (i.e. popping or cracking noises) detected during 58 cervicothoracic HVLA thrust manipulation procedures. Duration of the thrust procedure Similar to Dunning et al,26 but unlike three previous studies,23,55,56 the time interval between the beginning of first cavitation and the end of the last cavitation was used to represent the duration of the actual thrusting procedure from onset to arrest in the current study; nevertheless, the mean duration of a single cervicothoracic HVLA thrust manipulation was found to be 60.77 ms (95%CI 28.25, 97.42), a value that is slightly shorter but still consistent with Triano56 (135 ms), Herzog et al23 (80-100 ms), Ngan et al55 (158 ms) and Dunning et al26 (97 ms). Notably, Triano56 measured the duration of the thrusting procedure by analyzing force-time history graphs for a

C2-3 lateral break manipulation; whereas, Ngan et al55 used a four camera motion analysis system to measure head on trunk angular movements (and indirectly thrust duration) during lower cervical rotational manipulations in eight asymptomatic subjects. Additionally, Herzog et al23 measured thrust duration using “instantaneous acceleration signals” from a mechanical accelerometer during T4 posterior to anterior thrust manipulations in 28 subjects with thoracic pain. Therefore, considering the different instrumentation and analytical methods used in each of the previous studies,23,26,55,56 there does not appear to be a consistent reference standard for measuring thrust duration. Nevertheless, to date, this study is the first to report the thrust duration for a manipulation technique that targets the T1-2 articulations. Clinical relevance of the cavitation sounds The cavitation sound is traditionally considered by many physical therapists, chiropractors, and osteopaths to be an important indicator for the successful technical delivery of an HVLA thrust manipulation.19,20,22,23,25,35,39,40,56 However, four previous studies31-34 have suggested that the “audible pop” following HVLA thrust manipulation is not related to the clinical outcomes of pain and/or disability. Nevertheless, these authors31-34 investigated the thoracic and lumbopelvic regions, not the cervical spine or CTJ. Notably, many clinicians19,22,35 and researchers20,21,36-42 still appear to repeat the HVLA thrust manipulation if they do not hear or palpate popping sounds. Moreover, Evans and Lucas27 proposed the “audible popping”, or the “mechanical response” that “occurs within the recipient”, should be present to satisfy the criteria for a valid manipulation.27 Understanding whether the cavitation phenomenon during cervicothoracic HVLA thrust manipulation is an ipsilateral, contralateral or bilateral event will help inform practitioners of spinal manipulative therapy in selecting the appropriate technique that will most effectively target the dysfunctional articulation with the ultimate goal of reducing pain and disability. More specifically, considering the findings of previous studies17,20,25 and based on the results of our study, in order to maximize the likelihood that the target articulation is indeed manipulated, the practitioner should stand on the target side of the

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CTJ when performing a â&#x20AC;&#x153;lateral breakâ&#x20AC;? cervicothoracic HVLA thrust manipulation with the patient in prone, i.e., the short lever applicator of the practitioner should be placed on the side opposite the target or symptomatic articulation. Limitations It should be recognized that the morphology and arthrokinematics of the zygapophyseal joints are distinct to this region; thus, the results should be extrapolated to other spinal regions with caution. Furthermore, the results of this study cannot be generalized to cervicothoracic manipulation techniques that use different combinations of primary, secondary, physiologic and accessory component levers. One further limitation of this study is that only one practitioner administered all of the cervicothoracic thrust manipulations; while this enhances internal validity it also compromises generalizability. Future research should determine the vertebral levels at which the cavitation sounds are emanating from and investigate the clinical significance of the cavitation phenomenon following cervicothoracic HVLA thrust manipulation to determine whether a relationship exists between the number of cavitations and change in the clinical outcomes of pain and disability in various subgroups of patients. CONCLUSIONS Cavitation was significantly more likely to occur unilaterally, and on the side contralateral to the shortlever applicator contact, during cervicothoracic HVLA thrust manipulation. Most subjects produced three to five cavitations (i.e. popping or cracking noises) during a single lateral break HVLA thrust manipulation targeting the right or left T1-2 articulation; therefore, practitioners of spinal manipulative therapy should expect multiple cavitation sounds when performing HVLA thrust manipulation to the CTJ. Furthermore, the traditional manual therapy approach of targeting a single ipsilateral or contralateral facet joint during the delivery HVLA thrust manipulation may not be realistic. Whether the multiple cavitation sounds found in this study emanated from the same joint, adjacent ipsilateral or contralateral facet or uncovertebral joints, or even extra-articular soft-tissues remains to be elucidated. Due to the presence of multi-peak energy bursts and sounds of

multiple frequencies, neither the cavitation hypothesis (i.e. intra-articular gas bubble collapse) nor the tribonucleation hypothesis (i.e. cavity inception within synovial fluid) alone appear able to explain all of the audible sounds during HVLA thrust manipulation, and the possibility remains that several phenomena may be occurring simultaneously. REFERENCES 1. Fernandez-de-Las-Penas C, Alonso-Blanco C, Cleland JA, et al. Changes in pressure pain thresholds over C5-C6 zygapophyseal joint after a cervicothoracic junction manipulation in healthy subjects. J Manipulative Physiol Ther. 2008;31(5):332-337. 2. Salom-Moreno J, Ortega-Santiago R, Cleland JA, et al. Immediate changes in neck pain intensity and widespread pressure pain sensitivity in patients with bilateral chronic mechanical neck pain: a randomized controlled trial of thoracic thrust manipulation vs non-thrust mobilization. J Manipulative Physiol Ther. 2014;37(5):312-319. 3. Dunning JR, Cleland JA, Waldrop MA, et al. Upper cervical and upper thoracic thrust manipulation versus nonthrust mobilization in patients with mechanical neck pain: a multicenter randomized clinical trial. J Orthop Sports Phys Ther. 2012;42(1): 5-18. 4. Martinez-Segura R, De-la-Llave-Rincon AI, OrtegaSantiago R, et al. Immediate changes in widespread pressure pain sensitivity, neck pain, and cervical range of motion after cervical or thoracic thrust manipulation in patients with bilateral chronic mechanical neck pain: a randomized clinical trial. J Orthop Sports Phys Ther. 2012;42(9):806-814. 5. Masaracchio M, Cleland JA, Hellman M, Hagins M. Short-term combined effects of thoracic spine thrust manipulation and cervical spine nonthrust manipulation in individuals with mechanical neck pain: a randomized clinical trial. J Orthop Sports Phys Ther. 2013;43(3):118-127. 6. Salvatori R, Rowe RH, Osborne R, Beneciuk JM. Use of thoracic spine thrust manipulation for neck pain and headache in a patient following multiple-level anterior cervical discectomy and fusion: a case report. J Orthop Sports Phys Ther. 2014;44(6):440-449. 7. Cleland JA, Glynn P, Whitman JM, et al. Short-term effects of thrust versus nonthrust mobilization/ manipulation directed at the thoracic spine in patients with neck pain: a randomized clinical trial. Phys Ther. 2007;87(4):431-440. 8. Cross KM, Kuenze C, Grindstaff TL, Hertel J. Thoracic spine thrust manipulation improves pain, range of motion, and self-reported function in

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lumbar spinal manipulation. J Manipulative Physiol Ther. 2011;34(9):572-583. Dunning JR, Cleland JA, Waldrop MA, et al. Upper cervical and upper thoracic thrust manipulation versus nonthrust mobilization in patients with mechanical neck pain: a multicenter randomized clinical trial. J Orthop Sports Phys Ther. 2012;42(1): 5-18. Gibbons P, Tehan P. Manipulation of the spine, thorax and pelvis : an osteopathic perspective. 3rd ed. Edinburgh ; New York: Churchill Livingstone/ Elsevier; 2010. Herzog W, Zhang YT, Conway PJ, Kawchuk GN. Cavitation sounds during spinal manipulative treatments. J Manipulative Physiol Ther. 1993;16(8):523-526. Reggars JW, Pollard HP. Analysis of zygapophyseal joint cracking during chiropractic manipulation. J Manipulative Physiol Ther. 1995;18(2):65-71. Ross JK, Bereznick DE, McGill SM. Determining cavitation location during lumbar and thoracic spinal manipulation: is spinal manipulation accurate and speciﬁc? Spine (Phila Pa 1976). 2004;29(13):1452-1457. Dunning J, Mourad F, Barbero M, et al. Bilateral and multiple cavitation sounds during upper cervical thrust manipulation. BMC Musculoskelet Disord. 2013;14:24. Evans DW, Lucas N. What is ‘manipulation’? A reappraisal. Man Ther. 2010;15(3):286-291. Kawchuk G. The physics of spinal manipulation. Part 1. The myth of F = ma. J Manipulative Physiol Ther. 1992;15(3):212-213. Kawchuk GN, Fryer J, Jaremko JL, et al. Real-time visualization of joint cavitation. PLoS One. 2015;10(4):e0119470. Meal GM, Scott RA. Analysis of the joint crack by simultaneous recording of sound and tension. J Manipulative Physiol Ther. 1986;9(3):189-195. Bialosky JE, Bishop MD, Robinson ME, George SZ. The relationship of the audible pop to hypoalgesia associated with high-velocity, low-amplitude thrust manipulation: a secondary analysis of an experimental study in pain-free participants. J Manipulative Physiol Ther. 2010;33:117-124. Cleland JA, Flynn T, Childs JD, Eberhart SL. The audible pop from thoracic spine thrust manipulation and its relation to short-term outcomes in patients with neck pain. Journal of Manual and Manipulative Therapy. 2007;15:143-154. Flynn TW, Childs JD, Fritz JM. The audible pop from high-velocity thrust manipulation and outcome in individuals with low back pain. J Manipulative Physiol Ther. 2006;29(1):40-45.

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34. Flynn TW, Fritz JM, Wainner RS, Whitman JM. The audible pop is not necessary for successful spinal high-velocity thrust manipulation in individuals with low back pain. Arch Phys Med Rehabil. 2003;84:10571060. 35. Byfield D. Technique skills in chiropractic. Edinburgh: Churchill Livingstone; 2012. 36. Childs JD, Fritz JM, Flynn TW, et al. A clinical prediction rule to identify patients with low back pain most likely to benefit from spinal manipulation: a validation study. Ann Intern Med. 2004;141(12):920-928. 37. Cleland JA, Glynn P, Whitman JM, et al. Short-term effects of thrust versus nonthrust mobilization/ manipulation directed at the thoracic spine in patients with neck pain: a randomized clinical trial. Phys Ther. 2007;87(4):431-440. 38. Cross KM, Kuenze C, Grindstaff TL, Hertel J. Thoracic spine thrust manipulation improves pain, range of motion, and self-reported function in patients with mechanical neck pain: a systematic review. J Orthop Sports Phys Ther. 2011;41(9):633-642. 39. Gonzalez-Iglesias J, Fernandez-de-las-Penas C, Cleland JA, et al. Inclusion of thoracic spine thrust manipulation into an electro-therapy/thermal program for the management of patients with acute mechanical neck pain: a randomized clinical trial. Man Ther. 2009;14(3):306-313. 40. Gonzalez-Iglesias J, Fernandez-de-las-Penas C, Cleland JA, et al. Thoracic spine manipulation for the management of patients with neck pain: a randomized clinical trial. J Orthop Sports Phys Ther. 2009;39(1):20-27. 41. Puentedura EJ, Cleland JA, Landers MR, et al. Development of a clinical prediction rule to identify patients with neck pain likely to benefit from thrust joint manipulation to the cervical spine. J Orthop Sports Phys Ther. 2012;42(7):577-592. 42. Ruiz-Saez M, Fernandez-de-las-Penas C, Blanco CR, et al. Changes in pressure pain sensitivity in latent myofascial trigger points in the upper trapezius muscle after a cervical spine manipulation in pain-free subjects. J Manipulative Physiol Ther. 2007;30(8):578-583. 43. Kaltenborn FM. Mobilization of the spine. First ed. Oslo, Norway: Olaf Norlis Bokhandel; 1970. 44. Maitland GD. Vertebral manipulation. 2nd ed. London,: Butterworths; 1968. 45. Reggars JW. The manipulative crack. Frequency analysis. Australas Chiropr Osteopathy. 1996;5(2):39-44. 46. Cascioli V, Corr P, Till Ag AG. An investigation into the production of intra-articular gas bubbles and increase in joint space in the zygapophyseal joints of the cervical spine in asymptomatic subjects after

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IJSPT

CASE REPORT

RETURN TO RUNNING FOLLOWING A KNEE DISARTICULATION AMPUTATION: A CASE REPORT Angela R. Diebal-Lee1 Robert S. Kuenzi2 Christopher A. Rรกbago2,3

ABSTRACT Background and Purpose: The evolution of running-specific prostheses has empowered athletes with lower extremity amputations to run farther and faster than previously thought possible; but running with proper mechanics is still paramount to an injury-free, active lifestyle. The purpose of this case report was to describe the successful alteration of intact limb mechanics from a Rearfoot Striking (RFS) to a Non-Rearfoot Striking (NRFS) pattern in an individual with a knee disarticulation amputation, the associated reduction in Average Vertical Loading Rate (AVLR), and the improvement in functional performance following the intervention. Case description: A 30 year-old male with a traumatic right knee disarticulation amputation reported complaints of residual limb pain with running distances greater than 5 km, limiting his ability to train toward his goal of participating in triathlons. Qualitative assessment of his running mechanics revealed a RFS pattern with his intact limb and a NRFS pattern with his prosthetic limb. A full body kinematic and kinetic running analysis using 3D motion capture and force plates was performed. The average intact limb loading rate was four-times greater (112 body weights/s) than in his prosthetic limb which predisposed him to possible injury. He underwent a three week running intervention with a certified running specialist to learn a NRFS pattern with his intact limb. Outcomes: Immediately following the running intervention, he was able to run distances of over 10 km without pain. On a two-mile fitness test, he decreased his run time from 19:54 min to 15:14 min. Additionally, the intact limb loading rate was dramatically reduced to 27 body weights/s, nearly identical to the prosthetic limb (24 body weights/s). Discussion: This case report outlines a detailed return to run program that targets proprioceptive and neuromuscular components, injury prevention, and specificity of training strategies. The outcomes of this case report are promising as they may spur additional research toward understanding how to eliminate potential injury risk factors associated with running after limb loss. Keywords: Amputation, running, limb loading rate Level of Evidence: 4

Physical Therapist Assistant Program, Medical Education and Training Campus, JBSA Fort Sam Houston, Texas, United States of America 2 Center for the Intrepid, Department of Rehabilitation Medicine, Brooke Army Medical Center, JBSA Fort Sam Houston, Texas, United States of America 3 The Extremity Trauma and Amputation Center of Excellence, JBSA Fort Sam Houston, Texas, United States of America

CORRESPONDING AUTHOR Angela R. Diebal-Lee E-mail: angela.r.diebal.mil@mail.mil

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BACKGROUND A common goal following a lower extremity amputation is for the individual to return to their preinjury lifestyle. For those with active lifestyles and occupations, running can be central to achieving these endeavors. The development and evolution of running-specific prostheses has empowered recreational runners1 and Paralympic athletes to run farther and faster than previously thought possible.2-4 Despite the advancement of running prostheses,5,6 proper and symmetric mechanics are paramount to successful running and injury avoidance.7-9 Differing foot strike patterns between the prosthetic and intact limb is a common problem for individuals with unilateral lower extremity amputation, making a return to injury-free running challenging. Contemporary energy-storing-and-returning prosthetic limbs are optimized for running and require the runner to strike the ground in using a non-rear foot striking (NRFS) pattern. However, approximately 80% of runners use a rear foot striking (RFS) pattern;10,11 thus, an individual following a unilateral lower extremity amputation may attempt to run with asymmetric foot striking patterns. In able-bodied individuals, running with a RFS as opposed to a NRFS pattern significantly increases average vertical ground reaction force loading rates (AVLR), maximum braking forces, negative work of the ankle dorsiflexors, and negative work of the knee extensors.12,13 These biomechanical characteristics are associated with increased musculoskeletal injury risk through increased tissue stresses.14-19 Most healthy RFS runners demonstrate AVLRs between 60-70 body weights/second (BW/s) in laboratory studies.14 AVLRs greater than 70 BW/s are associated with tibial and metatarsal stress fractures, patellofemoral pain syndrome, and plantar fasciitis.9,11,20 Typical AVLRs for NRFS runners are observed from 40-60 BW/s; lower than the reported values that place people at risk for many common running injuries.11,12 An individual with a unilateral lower extremity amputation may be at similar risk for injuries if using a RFS pattern on their intact limb. In a case report of a 27 year-old (yo) with a unilateral transtibial amputation, average AVLR in the intact limb was 74 BW/s when using a RFS pattern.21 Based on the

abled-bodied literature, this high AVLR may predispose these runners to micro-traumas, stress injuries, or osteoarthritis in the intact limb.8,19 Further, the use of different foot strike patterns between prosthetic and intact limbs may lead to musculoskeletal imbalances and the development of secondary overuse injuries.19 In order to lessen these risks, rehabilitation programs should include instruction to foster foot strike symmetry between the prosthetic and intact limbs. Since the prosthetic limb cannot be changed to a RFS pattern, individuals must learn to produce a NRFS pattern with the intact limb. In the case report of the 27 yo with a unilateral transtibial amputation, the clinical investigators observed a 90% reduction of average AVLR to 39 BW/s in the intact limb when using a NRFS running pattern.21 These investigators attributed this change to NRFS running instruction and training during rehabilitation. While these results are promising, there is no evidence to support whether a change to a NRFS running pattern in more proximal amputees would elicit similar positive results. The purpose of this case report was to describe the successful alteration of intact limb mechanics from a RFS to a NRFS pattern in an individual with a knee disarticulation amputation, the associated reduction in AVLR, and the improvement in functional performance following the intervention. CASE DESCRIPTION The service member (SM) was a 30 yo male who sustained a traumatic right knee disarticulation amputation after being hit by a truck while riding his motorcycle. He also sustained a mild traumatic brain injury and multiple right metacarpal fractures that were stabilized with internal fixation hardware. He was 72 weeks post-amputation, 58 weeks postindependent ambulation, and had started running approximately 24 weeks prior to presentation. SM provided informed consent for all clinical procedures and approved the use of his data in this case report. At the time of presentation, he was discharged from rehabilitation services with all injuries well healed and no planned medical procedures. SM had a height of 1.83 m, body mass of 75.7 kg without his prosthesis, and residual limb length of 50 cm from the ipsilateral greater trochanter.

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SM utilized prosthetic care services at the Center for the Intrepid, Brooke Army Medical Center (JBSA Fort Sam Houston, TX) for distribution, fitting, and maintenance of his prosthetic components. He used two separate prosthetic limbs; one for typical gait and one for running. His gait prosthesis (mass 3.2 kg) consisted of a silicone gel liner with a custom Seal-In® ring (Ossur); passive-suction, ischial-level, flexible inner socket with carbon-fiber frame including anterior and posterior fenestrations; Ossur Total Knee® 2100; and Ossur Variflex XC Rotate® energystoring-and-returning foot. When wearing his gait prosthesis and walking shoes, SM’s leg lengths were 98 cm bilaterally from ipsilateral greater trochanters to the floor. SM’s running prosthesis (mass 3.2 kg) consisted of the same socket design and liner; Ottobock 3S80 Sport Knee; and Ossur Flex-Run™ foot (category 5HI). When wearing his running prosthesis and a neutral-minimalist running shoe, SM’s leg lengths were 95 cm (intact limb) and 98 cm (prosthetic limb) from ipsilateral greater trochanters to the floor. SM stated that his goals were to continue on active duty, return to his pre-injury level of fitness, and compete in triathlons. Long distance running (>5 km) was central to achieving these goals. Prior to his transfer to JBSA Fort Sam Houston, run training at a previous duty station allowed him to complete distances up to 16 km. However, he complained of bilateral lower extremity pain and residual limb swelling at distances greater than 5 km. Recovery often took several days which further limited his frequency of training reducing his running volume to less than 10 km per week. A baseline 2D running analysis was performed using the Dartfish application on an Apple iPad Air 2 by a physical therapist (ARD) with a running specialty certification. Analysis revealed an asymmetric ground-foot striking pattern; RFS with the intact limb and NRFS with the prosthetic limb. CLINICAL IMPRESSION #1 Based on clinical indices and previous literature findings, an asymmetric ground-foot striking pattern by SM was likely contributing to his discomfort and pain when running. Further comprehensive assessment of SM’s running mechanics was warranted.

EXAMINATION A 3D biomechanical running analysis was performed at the Center for the Intrepid as SM ran at a self-selected velocity on a level 16 m runway. Motion capture collection and analysis procedures were similar to those previously reported.21 Full body kinematic data were collected at 120 Hz using a 26-camera infrared motion capture system (Motion Analysis Corporation, Santa Rosa, CA) to track 57 reflective markers in a modified HelenHayes marker set placed on hand, arm, head, trunk, pelvis, thigh, shank and foot segments. The pylon of the prosthetic limb was modeled as a shank segment and markers were attached to running foot as to follow its contour (Figure 1). Kinetic data were collected at 1200 Hz (AMTI, Inc., Watertown, MA). Temporal spatial, kinematic, and kinetic parameters were quantified using Visual 3D software (C-Motion Inc., Rockville, MD). A minimum of eight strides per limb were used in the analyses. Kinematic and kinetic data were time-normalized to 100% gait cycle. Peak values for each kinematic and kinetic parameter of interest were extracted using a custom MATLAB program (Mathworks Inc., Natick, MA). Average vertical loading rate (AVLR) was defined as the slope of the vertical ground reaction force curve between 20% and 80% of the time to first impact peak.22 AVLR was normalized to SM’s body weight plus the weight his running prosthesis. SM’s average self-selected running velocity was 3.066 ± 0.048 m/s (Table 1). He spent 16.8% more time in stance on his intact limb compared to his prosthetic limb. He also spent 8.8% less time in swing with his intact limb compared to his prosthetic limb. These temporal-spatial asymmetries corresponded to observed asymmetric ground-foot striking patterns. Sagittal ankle and knee angles of both limbs are presented in Figures 2 and 3, respectively. As typical when RFS, at initial contact; SM struck with greater dorsiflexion and knee flexion of the intact limb than the prosthetic limb (Table 2). In contrast, SM demonstrated a NRFS pattern with his prosthetic limb as indicated by an apparent plantarflexed prosthetic ankle at initial contact. RFS with the intact limb lead to in an early vertical ground reaction peak at approximately 4% of

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the gait cycle (Figure 4). The resulting AVLR in the intact limb was approximately four-times greater than in his prosthetic limb (Figure 5, 112 ± 37 vs. 28 ± 2 body weights/s). CLINICAL IMPRESSION #2 Temporal-spatial, kinematic, and kinetic asymmetries associated with asymmetric ground-foot striking likely limited SM’s running tolerance and efficiency. Further, high AVLRs in the intact limb could potentially lead to repetitive stress injuries if not reduced.9 Therefore, an intervention focused on SM learning an intact limb NRFS pattern was developed. INTERVENTION A physical therapist (ARD) administered a threeweek long running intervention consisting of five, forty-five minute, running sessions (Table 3). Running instruction was conducted on flat outdoor surfaces as well as on a basketball court in a nearby gymnasium. A critical component of the intervention was patient education for the purposes of injury prevention, self-awareness of running errors, and modification of running mechanics. SM was guided through an Army funded running video which focused on injury prevention techniques through the identification of common running errors and instruction on proper running mechanics.23 Frequent feedback for the purpose of self-awareness was reviewed with SM using a Dartfish video at the conclusion of the running instruction each day to help demonstrate and correct running errors.

Figure 1. Biomechanical model used to determine kinematics and kinetics. Red circles are the reﬂective markers placed on body segments shown as bones. The right knee is lower than the contralateral intact knee as the prosthetic knee is below the anatomical end of the distal femur. The pylon that attaches to the prosthetic knee and foot is modeled as a shank to provided knee kinematics. The distal portion of the running foot is modeled as a single segment to provide ankle kinematics relative to the shank. Reﬂective markers were placed along the entire contour of the prosthetic foot for future analysis related to a multi-axis joint.

Specific running drills and exercises were used to enforce patient education concepts in order to eliminate running errors and modify running technique. Training drills and exercises were performed

Table 1. Self-selected running velocity (m/s), stance times (s), and swing times (s) the intact and prosthetic limbs pre- and post-treatment (Tx). mean ± sd Condi on - Limb

PreTx - Intact

PreTx - Prosthesis

PostTx - Intact

PostTx - Prosthesis

Velocity

3.066 ± 0.048

---

3.079 ± 0.062

---

Stance me

0.257 ± 0.009

0.220 ± 0.004

Swing me

0.436 ± 0.017

0.478 ± 0.011

0.406± 0.012

0.407 ± 0.009

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Figure 2. Average intact and prosthetic limb ankle angles through the gait cycle prior to (PreTx) and following (PostTx) the running intervention. Dorsiﬂexion (DF) is shown in the positive Y-axis and plantarﬂexion (PF) in the negative.

Figure 4. Normalized vertical ground reaction forces in the intact and prosthetic limbs through the gait cycle prior to (PreTx) and following (PostTx) the running intervention. The ﬁrst peak was used to determine average vertical loading rates.

Figure 3. Average intact and prosthetic limb knee angles through the gait cycle prior to (PreTx) and following (PostTx) the running intervention. Knee ﬂexion (Flex) is shown in the positive Y-axis and extension (Ext) in the negative.

Figure 5. Average vertical loading rates in the intact and prosthetic limbs prior to (PreTx) and following (PostTx) the running intervention. The type of ground-foot striking pattern is indicated as rear foot striking (RFS) or non-rear foot striking (NRFS).

Table 2. Ankle and knee angles (degrees) at initial contact (IC) and peak knee angle during swing (Sw) for the intact and prosthetic limbs pre- and post-treatment (Tx). mean ± sd Condi on – Limb

PreTx - Intact

PreTx - Prosthesis

PostTx - Intact

PostTx - Prosthesis

Ankle dorsiﬂexion @ IC

9.8 ± 1.6

-2.3 ± 0.8

-8.7 ± 2.4

-1.5 ± 1.0

Knee ﬂexion @ IC

7.7 ± 1.4

-1.9 ± 0.2

16.7 ± 2.6

-1.5 ± 0.8

Peak knee ﬂexion in Sw

68.1 ± 3.7

84.1 ± 0.4

100.7 ± 8.5

83.6 ± 0.5

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Table 3. Five, forty-ďŹ ve-minute running sessions consisted of the following drills and time intervals.

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Table 3. Five, forty-ďŹ ve-minute running sessions consisted of the following drills and time intervals. (continued)

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Table 3. Five, forty-ďŹ ve-minute running sessions consisted of the following drills and time intervals. (continued)

as described in Table 3 and Figures 6-21, which focused on visualization of running technique, selfawareness of running errors, motor re-training, body weight perception, and dynamic postural alignment. Following each drill a 10-20 meter run was conducted to determine if changes to his running mechanics occurred with that specific drill. At the end of each session the patient was given three drills for his home exercise program. He was initially instructed to perform these three drills at regular intervals of 100 meters during his training runs. As he began to consistently demonstrate a NRFS running pattern he was instructed to increase running distances during his training runs. OUTCOMES An identical 3D biomechanical running analysis was performed post-intervention; four weeks following the initial analysis. SM successfully adopted a NRFS pattern in the intact limb as indicated by the

plantarflexed position of his intact ankle at initial contact (Figure 2 and Table 2). His new symmetric ground-foot striking pattern appeared to reduce temporal-spatial, kinematic, and kinetic asymmetries. SMâ&#x20AC;&#x2122;s self-selected running velocity was nearly identical across both running analysis sessions (Table 1). Intact limb stance time was reduced and identical to the prosthetic limb following the running intervention. Likewise, swing times were less than 0.5% different between limbs. AVLRs decreased by 76% and 14% in the intact and prosthetic limbs, respectively (Figures 4-5). The AVLRs were nearly identical with a 13% greater AVLR in the intact limb compared to prosthetic limb. The reduction in intact limb AVLR was likely associated with NRFS and an increase in intact limb knee flexion at initial contact (Table 2). With the improvements in running mechanics, SM chose to lengthen his prosthetic limb by 1 cm to allow for additional foot deformation, thus

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Figure 8. Body Weight Shifting/Body Weight Perception (Reprinted with permission from Pose Method) Figure 6. Running in Place (Reprinted with permission from Pose Method)

Figure 9. Wall Fall (Reprinted with permission from Pose Method)

Figure 7. Two leg hops (Reprinted with permission from Pose Method)

increasing energy storage and return. Substantial functional improvements were associated with the observed changes in running mechanics. The SMâ&#x20AC;&#x2122;s two-mile run time on the Army Physical Fitness Test improved from 19:54 mins to 15:14 mins, pre to

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Figure 10. Foot Tapping (Reprinted with permission from Pose Method)

Figure 13. Infantry Run (Reprinted with permission from Pose Method)

Figure 11. Skipping (Reprinted with permission from Pose Method)

Figure 14. Change of Support (Reprinted with permission from Pose Method)

post intervention. Pain with running decreased from 6/10 to 0/10, allowing him to compete symptom free in two triathlons within four weeks of the intervention and the Army 10 mile run six months later.

Figure 12. Front Lunge (Reprinted with permission from Pose Method)

DISCUSSION SM reported to the physical therapist (ARD) due to an inability to comfortably run distances greater than 5 km even after previous running instruction and training. This limitation restricted his ability to

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Figure 15. Running in Place (Reprinted with permission from Pose Method)

his prosthetic limb. At the conclusion of his training, 3D biomechanical running analyses revealed a successful adoption of a NRFS pattern which decreased the intact limb loading rate and temporal-spatial asymmetries. SM was also successful in competing in several running events; completing distances up to 10 miles.

Figure 16. Prone Hip Dips (Reprinted with permission from Pose Method)

comfortably compete in desired running events such as triathlons, military runs, and half marathons. 3D biomechanical running analyses revealed temporalspatial asymmetries between his intact and prosthetic limbs; to include a 300% higher loading rate in the intact limb, which we hypothesized was due to a RFS running pattern. SM participated in five run training sessions over a period of three weeks to learn a NRFS pattern with his intact limb to match

Increased loading rates are known to cause a greater number of running related musculoskeletal injuries in non-amputees; often related to a RFS running pattern. Individuals with limb loss who run with a RFS on the intact limb may also be susceptible to similar musculoskeletal injuries. Further, increased loading rates resulting from a RFS pattern may worsen conditions that amputees are predisposed for, such as skin breakdown, residual limb swelling, and an inability to tolerate higher level activities due to pain.24 In order to lessen the risk for certain types of musculoskeletal injuries during running implementation, rehabilitation programs should consider instruction to foster NRFS running patterns and decreased limb loading rates. Pre-intervention, SMâ&#x20AC;&#x2122;s intact limb AVLR was 300% greater than the prosthetic limb and above a 70 BW/s threshold reported to increase risk for tibial and metatarsal stress fractures, patellofemoral pain syndrome, and plantar fasciitis.9,14,20 Following our three week running intervention, intact limb AVLR decreased to 27 BW/s; which was a 315% decrease from pre-intervention values. Similar decreases in limb loading rate, following a four week

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Figure 17. Prone Resistance Pull (Reprinted with permission from Pose Method)

Figure 18. Supine Resistance Pull (Reprinted with permission from Pose Method)

NRFS running intervention, were reported in an individual with a unilateral transtibial amputation.21 Despite SMâ&#x20AC;&#x2122;s running instruction prior to the intervention described in this case report, he was unable to comfortably tolerate running distances greater than 5 km. He participated in distances up to 10 miles but he experienced significant pain and discomfort, which required several down days to recover. Even though SM was able to complete distances up to 10 miles, he did so with poor running mechanics and without the understanding of how to properly train for these long distance running events. Essentially he faced common roadblocks that many able-bodied runners encounter. History of previous injury, high weekly running mileage/volume, and increased impact force,7,25,26 can all contribute significantly to

musculoskeletal running injuries. In addition, learning to run with a prosthesis may be more challenging and result in poor running mechanics; and when left uncorrected may increase the risk for musculoskeletal injury. Individuals with unilateral limb loss are five to 10 times more likely to suffer secondary musculoskeletal changes in their intact limb compared to those without limb loss.27-30 Thus, compared to able-bodied runners, an individual with limb loss may be at greater risk for musculoskeletal injuries when increasing mileage with uncorrected running mechanics. Correcting poor running mechanics prior to increasing mileage is essential to preventing injuries. In this regard, a training program for able-bodied individuals and individuals with amputation are

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Figure 19. Running with Elastic Bands (Reprinted with permission from Pose Method)

Figure 20. Running with the EZ Run Belt (Reprinted with permission from Pose Method)

Figure 21. Eccentric Hamstring (Reprinted with permission from Pose Method)

comparable. The program should consist of multiaxial, neuromuscular, and proprioceptive activities to decrease injury risk, evenly distribute musculoskeletal stresses, and improve running economy.31,32 Additionally, a runnerâ&#x20AC;&#x2122;s cadence should be increased to 180 steps/min or more so the foot strikes directly under the runner, thereby eliminating the RFS. Increased cadence and a NRFS ground contact style

will decrease the magnitude of loading throughout the lower extremity and together are crucial to mitigating common running injuries.33,34 Depending on the type of running prosthesis used, kinematic running patterns may become asymmetric and/or deviate from proper running form. SM utilized an articulating prosthetic knee (3S80) which

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demonstrated kinematics similar to the intact knee (Figure 3). Following the intervention, the prosthetic knee kinematics did not appear to change but the intact knee flexion during flight improved with the change to an NRFS pattern. Some amputee runners utilize non-articulating knees or no knee at all (i.e. straight pylon attached to running foot). A nonarticulating knee can cause a hip circumduction of the prosthetic limb in order to clear the prosthetic foot during flight. The decision to utilize an articulating versus a non-articulating knee is often a decision made between the prosthetist, therapist, and the patient. Considerations to ponder when choosing an articulating or non-articulating knee are energy costs and running mechanics. In the study conducted by Highsmith et al,33 energy costs, rate of perceived exertion, and heart rate was reduced and faster maximal speed was reached with an articulating knee versus a non-articulating knee. Energy costs for individuals with a transfemoral amputation are significantly greater than non-amputees. It is, therefore, imperative to incorporate prosthetic devices that decrease energy costs and improve ambulatory performance; especially in high functioning athletes. After discussions, which included literature review and sharing of clinical experience, SM chose to run with an articulating knee. Since the intervention, he has been able to successfully utilize his articulating running prosthesis for distances up to 13 miles without detriment. Six months following intervention, he placed third in a 10-mile run and completed one half marathon, both without pain. Presently, he is in the World Class Athlete Program training for a triathlon in the next Paralympics. CONCLUSION Currently, limited knowledge exists regarding specific training protocols for returning individuals with lower extremity limb loss to run. Additionally, it is not known whether these individuals require additional recovery time for tissue healing to prevent skin breakdown and other complications. This case report outlines a detailed return to run program, which addresses numerous proprioceptive and neuromuscular factors, targets injury prevention, and follows specificity of training strategies. Further study is warranted since the findings of this case report cannot be generalized to a larger population. The

outcomes of this case report are promising as they may spur additional research toward understanding how to eliminate potential injury risk factors associated with running following limb loss. The training regimen described in this case study will hopefully allow individuals with limb loss to improve their quality of life through a return to running sports. REFERENCES 1. Brown MB, Millard-Stafford ML, Allison AR. Running-speciďŹ c prostheses permit energy cost similar to nonamputees. Med Sci Sports Exerc. 2009;41(5):1080-1087. 2. Burkett B. Technology in Paralympic sport: performance enhancement or essential for performance? Br J Sports Med. 2010;44(3):215-220. 3. Hsu MJ, Nielsen DH, Yack HJ, Shurr DG. Physiological measurements of walking and running in people with transtibial amputations with 3 different prostheses. J Orthop Sports Phys Ther. 1999;29(9):526-533. 4. Pepper M, Willick S. Maximizing physical activity in athletes with amputations. Curr Sports Med Rep. 2009;8(6):339-344. 5. Shultz AH, Lawson BE, Goldfarb M. Running with a powered knee and ankle prosthesis. IEEE Trans Neural Syst Rehabil Eng. 2015;23(3):403-412. 6. Nolan L. Carbon ďŹ bre prostheses and running in amputees: a review. Foot Ankle Surg. 2008;14(3):125129. 7. Yeung EW, Yeung SS. A systematic review of interventions to prevent lower limb soft tissue running injuries. Br J Sports Med. 2001;35(6):383-389. 8. Hreljac A. Impact and overuse injuries in runners. Med Sci Sports Exerc. 2004;36(5):845-849. 9. Zifchock RA, Davis I, Higginson J, McCaw S, Royer T. Side-to-side differences in overuse running injury susceptibility: a retrospective study. Hum Mov Sci. 2008;27(6):888-902. 10. Hasegawa H, Yamauchi T, Kraemew WJ. Foot strike patterns of runners at the 15-km point during an elite-level half marathon. J Strength Cond Res. 2007;21(3):888-893. 11. Lieberman DE, Venkadesan M, Werbel WA, et al. Foot strike patterns and collision forces in habitually barefoot versus shod runners. Nature. 2010;463(7280):531-535. 12. Arendse RE, Noakes TD, Azevedo LB, Romanov N, Schwellnus MP, Fletcher G. Reduced eccentric loading of the knee with the pose running method. Med Sci Sports Exerc. 2004;36(2):272-277.

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13. Goss DL, Gross MT. A comparison of negative joint work and vertical ground reaction force loading rates in Chi runners and rearfoot-striking runners. J Orthop Sports Phys Ther. 2013;43(10):685-692. 14. Davis IS, Bowser BJ, Hamill J. Vertical Impact Loading in Runners with a History of Patellofemoral Pain Syndrome. Med Sci Sports Exerc. 2010;42(5):682-682. 15. Davis IS, Bowser BJ, Mullineaux DR. Greater vertical impact loading in female runners with medically diagnosed injuries: a prospective investigation. Br J Sports Med. 2016;50(14):887-892. 16. Diebal AR, Gregory R, Alitz C, Gerber JP. Forefoot running improves pain and disability associated with chronic exertional compartment syndrome. Am J Sports Med. 2012;40(5):1060-1067. 17. Pohl MB, Hamill J, Davis IS. Biomechanical and anatomic factors associated with a history of plantar fasciitis in female runners. Clin J Sport Med. 2009;19(5):372-376. 18. Zadpoor AA, Nikooyan AA. The relationship between lower-extremity stress fractures and the ground reaction force: a systematic review. Clin Biomech (Bristol, Avon). 2011;26(1):23-28. 19. Zifchock RA, Davis I, Hamill J. Kinetic asymmetry in female runners with and without retrospective tibial stress fractures. J Biomech. 2006;39(15):2792-2797. 20. Zadpoor AA, Nikooyan AA. Modeling muscle activity to study the effects of footwear on the impact forces and vibrations of the human body during running. J Biomech. 2010;43(2):186-193. 21. Waetjen L, Parker M, Wilken JM. The Effects of Altering Initial Ground Contact in the Running Gait of a Transtibial Amputee: A Case Report. Prosthet Orthot Int. 2012;36(3):356-360. 22. Milner CE, Ferber R, Pollard CD, Hamill J, Davis IS. Biomechanical factors associated with tibial stress fracture in female runners. Med Sci Sports Exerc. 2006;38(2):323-328. 23. Diebal AR. Pose Running. YouTube. Jan 6, 2017, 2017. 24. Gailey R, Allen K, Castles J, Kucharik J, Roeder M. Review of secondary physical conditions associated with lower-limb amputation and long-term prosthesis use. J Rehabil Res Dev. 2008;45(1):15-29.

25. Bullock SH, Jones BH, Gilchrist J, Marshall SW. Prevention of physical training-related injuries recommendations for the military and other active populations based on expedited systematic reviews. Am J Prev Med. 2010;38(1 Suppl):S156-181. 26. Hreljac A, Marshall RN, Hume PA. Evaluation of lower extremity overuse injury potential in runners. Med Sci Sports Exerc. 2000;32(9):1635-1641. 27. Norvell DC, Czerniecki JM, Reiber GE, Maynard C, Pecoraro JA, Weiss NS. The prevalence of knee pain and symptomatic knee osteoarthritis among veteran traumatic amputees and nonamputees. Arch Phys Med Rehabil. 2005;86(3):487-493. 28. Pruziner AL, Werner KM, Copple TJ, Hendershot BD, Wolf EJ. Does Intact Limb Loading Differ in Servicemembers With Traumatic Lower Limb Loss? Clin Orthop. 2014;472(10):3068-3075. 29. Robbins CB, Vreeman DJ, Sothmann MS, Wilson SL, Oldridge NB. A review of the long-term health outcomes associated with war-related amputation. Mil Med. 2009;174(6):588-592. 30. Struyf PA, van Heugten CM, Hitters MW, Smeets RJ. The prevalence of osteoarthritis of the intact hip and knee among traumatic leg amputees. Arch Phys Med Rehabil. 2009;90(3):440-446. 31. Heiderscheit BC, Chumanov ES, Michalski MP, Wille CM, Ryan MB. Effects of step rate manipulation on joint mechanics during running. Med Sci Sports Exerc. 2011;43(2):296-302. 32. Midgley AW, McNaughton LR, Jones AM. Training to enhance the physiological determinants of longdistance running performance: can valid recommendations be given to runners and coaches based on current scientiďŹ c knowledge? Sports Med. 2007;37(10):857-880. 33. Highsmith MJ, Kahle JT, Miro RM, Mengelkoch LJ. Bioenergetic Differences during Walking and Running in Transfemoral Amputee Runners Using Articulating and Non-Articulating Knee Prostheses. Technol Innov. 2016;18(2-3):159-165. 34. Rothschild C. Running Barefoot or in Minimalist Shoes: Evidence or Conjecture? Strength Cond J. 2012;34(2):8-17.

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IJSPT

CASE REPORT

THE EFFECTS OF A MULTIMODAL REHABILITATION PROGRAM ON PAIN, KINESIOPHOBIA AND FUNCTION IN A RUNNER WITH PATELLOFEMORAL PAIN Samuele Passigli, PT1 Pietro Capacci, PT2 Emanuele Volpi, PT1

ABSTRACT Background and Purpose: Multimodal interventions possess the strongest evidence in the long-term management of patellofemoral pain, but despite receiving evidence-based treatments that are initially effective many patients report recurrent or persistent symptoms for years after the initial diagnosis. Untreated psychological factors could be a possible explanation for persistent symptoms and poor treatment outcome. The purpose of this case report was to describe and evaluate the effects of a multimodal rehabilitation program that included pain education, a graded program of lower limb strengthening, and running retraining on pain, kinesiophobia, and function in a runner with patellofemoral pain. Case description: The subject was a 37-year-old female runner reporting a 12-month history of anterior knee pain with previous failed physiotherapeutic treatment. She discontinued running when symptoms gradually worsened, approximately six months after initial onset. She was advised to avoid painful activities. Clinical examination revealed pain during the performance of a weight-bearing functional task, fear of movement, and functional limitations. Treatment focused on pain education, self-management strategies, and progressive loading of the involved tissues through a graduated program of exercises and running retraining. Outcomes: Clinically meaningful improvements were seen in pain, kinesiophobia, and function following a 21-week multimodal rehabilitation program. Discussion: This case report illustrates several important aspects of clinical reasoning contributing to successful outcomes for a runner with patellofemoral pain. The multimodal rehabilitation program utilized was based upon the neurophysiology of pain (pain education) rather than the tissue pathology model. The findings from this case report may be used to benefit clinicians with similar subject presentations and drive future research into the use of these interventions based upon neurophysiology models of pain in the treatment of patellofemoral pain. Level of Evidence: Level 4 Key words: Kinesiophobia, pain education, patellofemoral pain, running

1 2

University of Rome, “Tor Vergata”, Rome, Italy Centro Fisioterapico FisioWorld, Bibbiena, Arezzo, Italy

CORRESPONDING AUTHOR Samuele Passigli, PT via Nazionale 33, 52018 Castel San Niccolò, Arezzo, Italy +393336006925 E–mail: samuelepas@gmail.com

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BACKGROUND AND PURPOSE Patellofemoral pain (PFP) is a prevalent multifactorial knee condition, accounting for 16.5% of all consultations in sports medicine clinics.1 Its impact is profound, often reducing the ability of those with PFP to perform sporting, physical activity and work-related activities pain-free. Multimodal interventions, which commonly consist of combined quadriceps and gluteal strengthening, along with appropriate patient education, possess the strongest evidence in the long term management of PFP.2 Furthermore, running gait retraining has been shown to be effective in reducing pain and improving function in runners with PFP.3 Nevertheless, despite receiving evidence-based treatments that are initially effective, the long term prognosis for PFP is still poor, with many patients reporting recurrent or persistent symptoms for years after the initial diagnosis.4 Growing evidence suggests that psychological features such as pain-related fear, anxiety, depression, catastrophising, and self-efficacy play a role in the development and maintenance of persistent musculoskeletal pain.5 They have also been associated with pain and disability and identified as barriers to recovery and as factors that limit the potential for improvement with rehabilitation. As a persistent musculoskeletal condition that is no longer considered self-limiting, PFP may also be characterised by the coexistence of physical and non-physical features.6,7 Psychological features such as fear-avoidance beliefs, anxiety, kinesiophobia and catastrophising are known to be prevalent and predictive of outcome in patients with PFP. Untreated psychological factors could be a possible explanation for persistent symptoms and poor treatment outcome. Therefore, co-interventions to reduce catastrophising beliefs and kinesiophobia may enhance the results. A rationale for an assessment and treatment approach that moves the focus away from a traditional biomedical and structural model of pain, towards one directed at the neurophysiology of pain, has been recently suggested.8 The purpose of this case report was to describe and evaluate the effects of a multimodal rehabilitation program that included pain education, a graded program of lower limb strengthening exercises, and running retraining on pain, kinesiophobia, and function in a runner with PFP.

CASE DESCRIPTION The subject in this study was a 37-year-old female reporting a 12-months history of right anterior knee pain, with an insidious onset. She had been a recreational runner for fifteen years and participated in road races including half marathons and marathons. Without symptoms, the subject ran approximately 40 km per week on paved roads. However, she discontinued running when symptoms gradually worsened, approximately six months after initial onset. The subject denied any trauma to the knee and any previous history of knee or lower limb problems. She reported no other health problems. The subject saw an orthopaedic surgeon five months prior, who diagnosed her with patellofemoral pain syndrome. Magnetic resonance imaging at that time revealed small patellar osteophytes and high signal intensity of the Hoffa fat pad. She was advised to avoid running, cycling, and painful activities. Swimming was suggested, which the subject had diligently performed ever since. Anti-inflammatory medication had been prescribed for this condition, but the subject found no relief with this intervention. Previous physiotherapy treatment involved two months of quadriceps strengthening exercises, patellar taping, physical modalities, and foot insoles, with very little benefit. The subject was very worried regarding the progressively worsening of symptoms and because the orthopaedic surgeon suggested an arthroscopic lateral retinacular release if the pain did not improve. Furthermore, the previous physiotherapist highlighted an abnormal patellar tracking as cause of her pain and the fact that this could lead to a premature wear of the cartilage if she did not avoid running and painful activities. The subject featured in this case report gave written informed consent to participate in the study and was informed that the data concerning the case report would be submitted for publication. OUTCOME MEASURES The outcome measures utilized in this case report included the Tampa Scale for Kinesiophobia-11 (TSK11), the Knee Outcome Survey â&#x20AC;&#x201C; Activities of Daily Living Scale questionnaire (KOS-ADLS), perceived knee pain during the performance of a forward stepdown test (FSD) measured using the Numeric Pain Rating Scale (NPRS), and self-reported posterolateral

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hip pain (NPRS). The outcome measures were collected at initial examination, after eight weeks, and again at 21 weeks. The TSK-11 is an 11-item questionnaire designed to assess fear of movement and reinjury while offering the advantage of brevity. All items are based on a 4-point Likert scale in which patient options range from strongly disagree to strongly agree. The TSK11 scores range from 11 to 44, with higher scores indicating a higher degree of kinesiophobia. Despite the shortened format, the TSK-11 has demonstrated similar factor structure, internal consistency, testretest reliability, and validity to the original TSK-17.9 The shortened version has been used extensively in orthopaedic populations, including patients with lower extremity disability.10 A minimally clinical important difference (MCID) in the level of kinesiophobia has been determined to be a 4-point difference in TSK-11 scores.9 The KOS-ADLS is a 14-item knee-specific questionnaire designed to evaluate symptoms and functional limitations experienced during activities of daily living in individuals with various knee disorders.11 Six items assess knee symptoms (pain, stiffness, swelling, instability, weakness, and limping) and eight items assess functional limitations (walking on a level surface, ascending and descending stairs, standing, kneeling, squatting, sitting, and rising from a chair). Each item is scored on a 6-point Likert scale (0â&#x20AC;&#x201C;5 points). Total score on 70 is converted to a 0â&#x20AC;&#x201C;100 point scale with 100 indicating the absence of symptoms and functional limitations. The KOS-ADLS has been validated in patients with PFP, with a reported MCID of 7 points.12 A systematic review pertaining to self-reported questionnaires used in individuals with PFP recommended the KOS-ADLS over other knee-specific scales based on its psychometric properties and clinical applicability for runners.13

10 (the worst imaginable pain). The NPRS has been validated in patients with PFP, with a reported MCID of 2 points.12 The NPRS was administered verbally. EXAMINATION The subject described two areas of pain. The first area was located over the anterior region of the right knee (labeled P1, Figure 1). Her highest pain intensity on the NPRS in this location was rated as 7/10. The subject described the pain as intermittent and acute, provoked while sitting for longer than 30 min with her knee in a flexed position, descending stairs, squatting, and walking more than 30 min. Symptom eased with rest. The second area of pain was located in the posterolateral region of the right hip (labeled P2, Figure 1). The intensity of pain was rated as 5/10. The subject described the pain as intermittent and deep, provoked while sitting for longer than 30 min with her knee in a flexed position and standing for a longer than two hours. Active and passive examination of the lumbar spine was performed to rule out spine contributions to her symptoms. Active examination tests for the lumbar

The FSD is a clinical test designed to replicate stair descent, a functional task often limited by PFP.14 This test has been shown to be significantly correlated with pain in patients with PFP, and has demonstrated high intrasubject reliability which is a prerequisite for detecting true changes after treatment.15,16 The 11-point NPRS is a self-administered measure of the intensity of pain, ranging from 0 (no pain) to

Figure 1. Pain diagram. P1 and P2 indicate two distinct areas of perceived pain by the subject, P1 = primary pain area, P2 = secondary pain area.

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spine consisted of active range of motion, followed by active range of motion with overpressure in flexion, extension, and sidebending and rotation. Passive testing was performed using both central and lateral posterior-to-anterior spring tests from the fifth lumbar to the tenth thoracic spinal segments. There was no reproduction of her buttock or knee symptoms with active or passive testing to the lumbar spine. The subject’s knee symptoms were not reproduced by the femoral slump test. Observation of the subject’s knee revealed no swelling, bruising, or obvious bony deformity. Observation did not show also any differences in lower limb muscle trophism or significant postural asymmetries. Palpation assessment of the knee showed tenderness over the medial aspect of the patella. No pain was noted on the patellar tendon or along knee medial or lateral joint lines. Soft tissue palpation revealed tender nodules in taut tissue band in the gluteus medius muscle, indicative of the likelihood of trigger points (TrPs) in the affected musculature. Deep palpation of gluteus medius muscle reproduced 7/10 pain in the P2 region. The subject’s range of motion for hip, knee and ankle, measurements were pain free and symmetrical bilaterally except for hip internal rotation with right internal rotation measured at 25 degrees and left internal rotation measured at 40 degrees. Performance of the FSD reproduced 7/10 pain over the anterior region of the knee. This functional task also showed excessive dynamic knee valgus. At this stage it was felt the examination had reached the maximum possible level of testing without exacerbating the subject symptoms, therefore the examination ended. The subject’s score on the KOS-ADLS was 49/100 and 30/44 on the TSK-11. Her goal was to reduce pain and return to running. CLINICAL IMPRESSION Subject report of anterior knee pain, aggravated by activities that load the patellofemoral joint (PFJ) during weight bearing, tenderness on patellar facet palpation, and pain on prolonged sitting were consistent with diagnosis of PFP.17 TrPs in the gluteus medius muscle were suspected as a potential source of pain in the posterolateral hip region. Subjects with PFP have a significant prevalence of TrPs in the

gluteus medius that may be an important variable to consider in the evaluation and treatment of this condition.18 The subject showed elevated signs of fear and anxiety of movement and physical activity. INTERVENTION The subject was seen for a total of 20 visits over the course of 20 weeks. First session Immediate pain relief should be a priority in order to gain subject trust, facilitate active engagement and optimize long term outcomes.2 Therefore, posterolateral hip pain (P2) was treated in the first session. According to the literature, the ability to definitively ascertain the exact location of a TrP is questionable, and examiner experience plays a positive role in determining the presence of a TrP.19,20,21 Identification of a tender nodule in a taut band of muscle along with reproduction of the subject’s subjective report of pain is the most clinically accurate way to recognize the presence of a TrP.22 Risks and potential complications were advised and written consent was obtained outlining common and serious adverse events associated with dry needling (DN) interventions. Common complications include bruising, vasovagal response, bleeding, and muscle soreness. More serious (but rare) complications include infection, broken needle, and pneumothorax.23 There were no reported contraindications to the use of DN. DN techniques are proposed to quickly reduce pain, remove peripheral sources of persistent nociceptive input, and improve range of motion in a host of pathological conditions, such as chronic lateral hip pain.24 DN was performed to the gluteus medius muscle at the area determined by deep palpation as a possible locations of the TrPs (Figure 2). The DN technique utilized ten fast-in/out movements in a cone pattern to attempt to target as many sensitive loci as possible within the tender nodule in the taut band of muscle. Appropriate patient education is essential to effective PFP management.2 Therefore, she was given the ‘Managing My Patellofemoral Pain’ education leaflet (created by Barton CJ and Rathleff MS) to facilitate knowledge retention and reduce emphasis on biomechanics.25 To facilitate adherence to home exercise and optimize recovery, she was given infor-

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Figure 2. Dry needling of the gluteus medius muscle.

mation on the importance of an active role over passive intervention and positive dose–response relationship between exercise and recovery.26 Intervention: Weeks 1-8 The subject was instructed to perform an exercise program divided into two phases. She was asked to perform prescribed exercises three times/ week. Weekly meetings were scheduled to ensure proper execution of exercises and gradual progression of loads. Exercises were dosed and progressed according to pain levels and number of repetitions reached. The subject used the pain-monitoring model to grade the level and dosage of exercise and physical activity.27 According to the pain-monitoring model, the pain was allowed to reach level 5 on the visual analog scale, where 0 is no pain and 10 is the worst pain imaginable, during and after the exercise training, but should have subsided by the following morning. She was educated that it was not injurious to feel some pain during exercise during the exercise training. Increased fear of painful movement may be negative for recovery if it causes the subject to not load enough during the exercise program. Allowing for some amount of pain may be necessary to ensure that the load is sufficient to create meaningful adaptive changes in the PFJ. Descriptions of biomedical models of pain (e.g. abnormal patellar tracking, pronated foot or limb malalignment) and anatomic findings (patellar osteophytes and high signal intensity of Hoffa fat pad) were actively discouraged and challenged with pain described as a

consequence of ‘de-conditioned’ tissue. A painful loaded exercise program focused on functional gains instead of symptom modification may have a key rule in fear avoidance education reconceptualizing pain, de-emphasizing anatomic findings and lower limb structural anomalies, and encouraging the patients to take an active role in their recovery.28,29 In Phase one floor exercises were performed to recruit gluteal, quadriceps, and core muscles (Figures 3-8).30,31 Repetitions were externally paced by a metronome. Externally paced resistance training may be used to introduce a skill based element that may improve motor control.32 The main focus of Phase two of the treatment program was to progressively increase the exercises functional demand and load on PFJ by progressing hip and quadriceps exercises to weight bearing (Figures 9,10). Abdominal planks were progressed through higher levels of difficulty (Figures 11,12). An external focus of attention (attention directed towards the outcome or effects of the movement) was also added by using simple strategies, such us a cup full of water (the subject was asked to perform exercises without spilling water) (Figure 13), or a SenMoCOR™ (Sensory Motor Control-Oriented Rehabilitation) System (the subject performed exercises with immediate visual feedback using laser and laser target) (Figure 14). An externally focused exercise training may enhance skill acquisition more efficiently, promote utilization of unconscious or automatic processes, increase compliance, and

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Figure 3. Side-lying hip abduction.

Figure 4. Side-lying clam with elastic band.

Figure 5. Supine two-leg bridge. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 675

Figure 6. Supine one-leg bridge.

Figure 7. Plank.

Figure 8. Side plank. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 676

increase the potential to transfer of improved motor skills to sport.33,34 Furthermore, the subject was advised to gradually start cycling with similar advice given on not to fear or avoid some amount of pain. In the authors’ clinical experience, cycling represents an effective exer-

cise strategy for improving strength of the kinetic chain and load capacity of PFJ. The subject reported complete pain relief in P2 and hip internal rotation range of motion was symmetrical bilaterally after the first session. Therefore, other DN sessions or treatments had not been performed for this symptom.

Figure 9. Squat with elastic band.

Interventions: Weeks 9-20 The subject was evaluated on an instrumented treadmill (MyRun, Technogym, FC, Italy) at a preferred running pace of 6 minutes/km. Foot strike pattern was determined visually by looking at the slow motion video recording. Variables of interest were cadence (steps per minute), ground contact time (the time the foot is in contact with the ground during each step), and vertical oscillation (the amount of “bounce” in running motion), collected using Garmin Fenix 3 watch and Garmin HRM-Tri strap (Garmin Ltd, Southampton, UK). The subject had a forefoot strike pattern. The subject’s running cadence was 168 steps per minute, ground contact time was 272 ms, and vertical oscillation was 9 cm. She reported 5/10 pain over the anterior region of the knee. The primary intervention was running gait retraining with step rate manipulation. To reduce PFJ contact forces and stress while running, she was asked to increase her step rate to 185 steps per minute. There is strong evidence indicating reduced PFJ stress/load with increasing the step rate by 10%.3,35,36 Step rate was maintained using an audible metronome set to 185 beats per minute. The subject was also provided with simple verbal cues

Figure 10. One-leg squat. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 677

Figure 11. One-leg plank.

Figure 12. Side plank with external focus.

that included “run quietly” and “reduce noise”. The subject was able to successfully achieve this cadence. After the step rate manipulation, the cadence was 185, ground contact time was 242 ms, and vertical oscillation was 7.7 cm. Pain improved to 3/10. The subject received a personalized running program provided over 12 weeks (three sessions per week with at least one rest day between running days). She was advised to maintain her new step rate during running practice with help of the metronome and to avoid downhill running. She was instructed to utilize constant auditory cueing from the metronome until she was able to easily and consistently match the cadence goal. The run time was gradually increased from five minutes to 30 minutes over the

12 sessions. No specific recommendation was provided regarding lower limb alignment during running. Weekly meetings were scheduled to ensure gradual progression of running training. She was not to exceed a 5/10 pain intensity during training, according to the pain-monitoring model. The subject was advised to continue the home exercise program and cycling as cross-training. OUTCOMES Outcomes are shown in Table 1. Over time, KOS-ADLS score improved from a 49/100 to 90/100 and TSK-11 from 30/44 to 16/44. Subject met the MCID of 7 points on the KOS-ADLS and 4 points on the TSK-11. Additionally, perceived knee pain during the performance of FSD decreased from 7/10 to 2/10, which was successful

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Figure 13. Step down with resistance from an elastic band and external focus.

in obtaining the required 2-point change to be a clinically meaningful improvement. The subject reported complete pain relief in the posterolateral region of the hip. Despite clinically relevant improvements with treatment, kinematics during the performance of FSD was unchanged, questioning the role of biomechanics as a contributor to pain in this subject. As previously described, factors thought to relate to PFP often remain after patientsâ&#x20AC;&#x2122; symptoms have resolved making their clinical importance difficult to determine.37 Subjectively, the subject reported that she was able to run without limitation. The subject also noted that her running pace had improved from 6:00 minutes/km to 5:45 minutes/km, but it is unclear whether this was related to the increased step rate, improved pain, reduced kinesiophobia, or factors related to training. An e-mail received nine months following discharge noted that the subject completed a marathon symptom free. DISCUSSION This case report provides several important aspects related to clinical reasoning contributing to the suc-

Figure 14. Squat with Sensory Motor Control-Oriented Rehabilitation.

Table 1. Outcome measures

cessful outcomes for a runner with PFP. There is no consensus on the pathogenesis of PFP, with numerous biomechanical factors attributed to the greater PFJ stress and the development of pain and disability.6 As a result, various conservative interventions based upon tissue pathology models of pain have been proposed, including hip and quadriceps strengthening, patellar taping, braces, foot orthoses,

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stretching, and soft tissue manipulation.2 PFP is not a self-limiting condition, with recurrent or persistent symptoms often remaining for years after the initial diagnosis.4 Furthermore, the presence of structural abnormalities of the PFJ is not associated with pain and disability in many patients.38 A possible reason for the continued pain is that PFP is a multifactorial condition and the treatments may not address all of the contributing factors in each patient. As in other musculoskeletal conditions in which chronic pain occurs, psychological and social factors may integrate these factors in order to explain and understand persistent or recurrent symptoms. Patients with long-lasting PFP show signs of kinesiophobia, pain catastrophising, fear-avoidance belief, anxiety, and excessively negative thoughts towards their pain and function.5,6,7 These factors were strong predictors of pain intensity, function, and disability in PFP patients.6,7 Therefore, co-interventions to reduce catastrophizing thinking and kinesiophobia may improve the long-term outcomes. A loaded exercise program that allowed for some pain during exercise and running retraining, with pain described as ‘de-conditioned’ tissue, along with appropriate patient pain education and self-management strategies, may desensitize the central nervous system and address fear avoidance, kinesiophobia, and catastrophising beliefs.8,29 A pain-monitoring model may guide patients on how to cope with and respond to pain during and after exercises.27 This model, within a framework that suggests ‘hurt does not equal harm’, may encourage patients to take an active role in their recovery, improve adherence to the exercise program, identify an appropriate load during exercises and running, and shift the focus to functional gains with rehabilitation instead of pain. Progressive loading of the PFJ may reduce local hyperalgesia and may alter central pain processing in individuals with PFP.39,40 Therefore, progressive hip and quadriceps strengthening may improve a patient’s envelope of function by enhancing load tolerance of the PFJ. Whether patellar taping should be first-line treatment for PFP is unclear.30 Patellar taping was not used because a previous healthcare practitioner had already used it without benefit to the subject. In addition, patellar taping could reinforce the subject’s beliefs on a pathoanatomical source of pain, such as

an abnormal patellar tracking, and consequently result in more kinesiophobia. The subject tolerated the DN intervention to the hip region very well with no side effects reported following treatment. She reported a hypoalgesic effect following the DN intervention and only complained of minimal muscle soreness that lasted approximately two hours following treatment at the area of needle penetration. The use of DN demonstrated positive outcomes for reducing pain associated with lower quarter TrPs in the short-term.41 Immediate pain relief should be a priority to gain patient trust, facilitate active engagement, and optimize outcomes.2 There are a number of limitations associated with this case report. A number of pathologies may cause posterolateral hip pain worsened with prolonged standing and sitting and should be considered in the clinician’s differential diagnosis. This case report uses only a single subject, as is typical of a case report. This is an inherent limitation offering only results that relate to this single subject that cannot be generalized to larger populations. The same physiotherapists who performed the initial evaluation also performed the treatments and completed the final evaluation. In future studies, using a different and potentially blinded evaluator could minimize potential bias. As there is a lack of evidence describing the rehabilitation of psychological factors in subjects with PFP, as well as relationships with pain and physical function, additional systematic research is needed to determine the exact contribution of a painful loaded exercise program to the overall treatment approach provided to this population. The findings from this case report may be used to benefit clinicians with similar subject presentations and drive future research into the use of these interventions based upon neurophysiology models of pain and pain education in the treatment of PFP. CONCLUSION This case report describes the history, assessment and treatment of a runner with PFP who demonstrated clinically meaningful improvements in pain, kinesiophobia, and function following a 21-week multimodal rehabilitation program. This program was based upon the neurophysiology of pain rather

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than the tissue pathology model. Treatment focused on education and loading the tissues over many weeks through a graded program of loaded exercises and running retraining.

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25. Barton CJ, Rathleff MS. ‘Managing My Patellofemoral Pain’: the creation of an education leaﬂet for patients. BMJ Open Sport & Exercise Medicine. 2016;2:e000086. 26. Rathleff MS, Roos EM, Olesen JL, et al. Exercise during school hours when added to patient education improves outcome for 2 years in adolescent patellofemoral pain: a cluster randomised trial. Br J Sports Med. 2015;49(6):406-12. 27. Thomeé R. A comprehensive treatment approach for patellofemoral pain syndrome in young women. Phys Ther. 1997;77(12):1690-703. 28. Selhorst M, Rice W, Degenhart T, et al. Evaluation of a treatment algorithm for patients with patellofemoral pain syndrome: a pilot study. Int J Sports Phys Ther. 2015;10(2):178-88. 29. Littlewood C, Malliaras P, Bateman M, et al. The central nervous system - An additional consideration in ‘rotator cuff tendinopathy’ and a potential basis for understanding response to loaded therapeutic exercise. Man Ther. 2013;18(6):468-72. 30. Crossley KM, van Middelkoop M, Callaghan MJ, et al. 2016 Patellofemoral pain consensus statement from the 4th International Patellofemoral Pain Research Retreat, Manchester. Part 2: recommended physical interventions (exercise, taping, bracing, foot orthoses and combined interventions). Br J Sports Med. 2016;50(14):844-52. 31. Earl JE, Hoch AZ. A proximal strengthening program improves pain, function, and biomechanics in women with patellofemoral pain syndrome. Am J Sports Med. 2011;39(1):154-63. 32. Rio E, Kidgell D, Moseley GL, et al. Tendon neuroplastic training: changing the way we think about tendon rehabilitation: a narrative review. Br J Sports Med. 2016;50(4):209-15. 33. Benjaminse A, Gokeler A, Dowling AV, et al. Optimization of the anterior cruciate ligament injury

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38. van der Heijden RA, de Kanter JL, Bierma-Zeinstra SM, et al. Structural Abnormalities on Magnetic Resonance Imaging in Patients With Patellofemoral Pain: A Cross-sectional Case-Control Study. Am J Sports Med. 2016;44(9):2339-46. 39. Pazzinatto MF, de Oliveira Silva D, Barton C, et al. Female Adults with Patellofemoral Pain Are Characterized by Widespread Hyperalgesia, Which Is Not Affected Immediately by Patellofemoral Joint Loading. Pain Med. 2016;17(10):1953-1961. 40. Østerås B, Østerås H, Torstensen TA, et al. Doseresponse effects of medical exercise therapy in patients with patellofemoral pain syndrome: a randomised controlled clinical trial. Physiotherapy. 2013;99(2):126-31. 41. Morihisa R, Eskew J, McNamara A, et al. Dry Needling in subject with muscular trigger points in the lower quarter: a systematic review. Int J Sports Phys Ther. 2016;11(1):1-14.

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IJSPT

CLINICAL COMMENTARY

A FOUR-PHASE PHYSICAL THERAPY REGIMEN FOR RETURNING ATHLETES TO SPORT FOLLOWING HIP ARTHROSCOPY FOR FEMOROACETABULAR IMPINGEMENT WITH ROUTINE CAPSULAR CLOSURE Benjamin D. Kuhns, MD, MS1 Alexander E. Weber, MD2 Brian Batko, BS2 Shane J. Nho, MD, MS2 Catherine Stegemann, PT2,3

ABSTRACT Hip preservation surgery has become more common over the past decade and is now a preferred treatment modality for an increasingly diverse array of pathology in the young, active patient with hip pain. In particular, hip arthroscopy has become an increasingly popular treatment choice for active patients diagnosed with femoroacetabular impingement (FAI). Appropriate postoperative rehabilitation is critical for overall patient success and optimal long-term outcome. As surgical techniques continue to evolve, rehabilitation protocols must adapt to accommodate changes in the surgical procedure and ultimately provide the safest and fastest recovery of function for the patient. One such surgical modification has been the incorporation of routine capsular closure as part of the treatment of FAI in the young, active patient. The purpose of this clinical commentary is to present a four-phase rehabilitation protocol for returning to sport following arthroscopic correction of FAI with routine capsular closure. Level of Evidence: 5 Keywords: Femoroacetabular impingement, four phase rehabilitation, hip arthroscopy, return-to-sport, routine capsular closure

Department of Orthopaedics & Rehabilitation, University of Rochester Medical Center, Rochester, NY, USA 2 Hip Preservation Center, Section of Young Adult Hip Surgery, Division of Sports Medicine, Department of Orthopedic Surgery, Rush Medical College of Rush University, Rush University Medical Center, Chicago, IL, USA 3 Department of Physical Therapy, Midwest Orthopaedics at Rush, Chicago, IL, USA

CORRESPONDING AUTHOR Benjamin D. Kuhns, MD, MS Department of Orthopaedics & Rehabilitation. University of Rochester Medical Center 601 Elmwood Avenue, Rochester, NY 14602 Box 665 E-mail: Benjamin_Kuhns@urmc.rochester.edu

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INTRODUCTION Femoroacetabular impingement (FAI) frequently presents in young and athletic patients, a population with a high expectations for return to sport and participation in recreational activities.1 While initially managed conservatively, many patients eventually require surgery, with many orthopedists choosing arthroscopic correction of the bony deformities underlying FAI and repair of the soft tissues.2–5 When indicated, hip arthroscopy for FAI is generally successful, with multiple authors reporting improved short and intermediate term outcome scores.6–9 However, in addition to successful intraoperative management, optimal outcomes are predicated on skillful and thorough postoperative rehabilitation.10 Arthroscopic techniques have substantially advanced over the last two decades. For example, with regard to acetabular labral treatment options, labral debridement was the only option in the early 2000s; however, current labral treatment options now include labral repair and reconstruction. Over the past fifteen years, rehabilitation protocols have been developed to parallel advances in arthroscopic techniques, with many individual protocols previously reported.10–20 In general, there are between four and five phases of rehabilitation, beginning immediately postoperatively, and extending to return to sport; however, the exact rehabilitation protocol varies between surgeons and surgical centers.20

Rehabilitation following hip arthroscopy must take into account the higher physical demands that make this patient population unique from the typically older patients undergoing total hip arthroplasty (THA) and subsequent rehabilitation.19 Additionally, the role of capsular management in hip arthroscopy has received increased scrutiny, with some groups reporting improved clinical outcomes following routine capsular closure..21,22 Given these findings, postoperative rehabilitation programs should be tailored to the individual to reflect evolving capsular management strategies. The purpose of this clinical commentary is to describe a four-phase physical therapy regimen for returning athletes to sport following hip arthroscopy with routine capsular closure for the treatment of FAI (Figure 1). SURGICAL PROCEDURE The senior surgeon author is a high volume hip arthroscopist who has previously described his surgical technique.23 Briefly, following appropriate in office work-up, diagnostic imaging, and failure of nonoperative treatment interventions including non steroidal anti-inflammatory medication, physical therapy, and corticosteroid injections, young, active patients with a diagnosis of FAI are indicated for arthroscopic treatment. Patients are placed in the supine position on a hip arthroscopy table with a well-padded perineal post (Figure 2A-B). Traction is applied and the central

Figure 1. 4-Phase return to sport following hip arthroscopy with routine capsular closure for the treatment of femoroacetabular impingement. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 684

compartment is accessed via the anterolateral and modified anterior portals (Figure 2C). An interportal capsulotomy (approximately 2-4 cm) is performed to expose the acetabular rim. Once in the central compartment a diagnostic arthroscopy is performed and the labrum, chondral surfaces, and acetabular rim are treated. After work in the central compartment is complete, traction is released and the peripheral compartment is accessed. All patients undergo T-capsulotomy through the distal anterolateral accessory (DALA) portal to assist with arthroscopic visualization of the femoral neck. A femoral osteochondroplasty is performed in the peripheral compartment to address cam-type FAI (Figure 2D). Dynamic fluoroscopic examination is used to confirm complete bony resection and absence of residual deformity (Figure 2E). At the conclusion of the case, the entire capsulotomy is repaired with high-strength mattress

Figure 2. Arthroscopic procedure for femoroacetabular impingement. Figure 2A: The patient is placed in the supine position with a well-padded perineal post. Figure 2B: The hip is distracted and the patient is prepared for surgery in sterile fashion. Figure 2C illustrates portal placement on a preoperative hip. Figure 2D demonstrates a cam deformity as seen through a T-capsulotomy. Figure 2E demonstrates a completed resection of the cam deformity. The completed capsular closure can be seen in Figure 2F.

sutures. The longitudinal portion of the T-capsulotomy is closed using three simple interrupted #2 high tensile strength sutures passed with a suture-shuttling device (Spectrum, Conmed Linvatec, Key Largo, FL). The interportal capsulotomy is subsequently closed with two simple interrupted #2 high tensile strength sutures in simple interrupted fashion using a capsular closure device (Injector, Stryker Sports Medicine, Greenwood Village, CO) (Figure 2F). The case is then completed with closure of the portals and the patient is transferred to the recovery with a brace designed to limit hip abduction and rotation.12 Following discharge, postoperative rehabilitation begins immediately. REHABILITATION A four phase rehabilitation protocol is followed to return patients to activity following hip arthroscopy with capsular closure. Criteria for advancement and common pitfalls are also presented. Phase One: Protect the Joint Phase one begins postoperatively with the goals of joint healing and achieving symmetrical range of motion by 6-8 weeks assuming no dysfunction of the contralateral nonsurgical lower extremity (Table 1). Initially, the patient is restricted to 20-pound footflat weightbearing. The restriction in weightbearing is related to reports of postoperative femoral neck fracture following femoral osteochondroplasty (CAM resection).24 Flat-foot weightbearing is preferred over toe-touch weightbearing due to concerns of Achilles contracture, hip flexor contracture, and lower extremity stiffness related to consistent ankle plantar flexion. Initial range of motion is limited to 90ยบ of flexion, 30ยบ of external rotation, 20ยบ of internal rotation with 90ยบ of hip flexion, and 30ยบ of abduction. Importantly, active open chain hip flexor activation should be avoided and extension should be limited to neutral, particularly in the presence of capsular repair.25 Additional precautions include no active straight leg raises and no sitting for longer than 30 minutes at a time. For the first physical therapy visit on postoperative day one, the upright stationary bike is started for 20 minutes with no resistance. The patient is trained on the usage of the brace and the continuous passive motion (CPM) device. Additionally, the patient and caregiver are

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Table 1. Phase One Rehabilitation

instructed on passive range of motion maneuvers including hip circumduction in neutral as well as 20ยบ to 30ยบ of flexion, and log rolls. Isometric exercises focusing on quadriceps, gluteal, and transversus abdominis (TA) activation are also emphasized. A home exercise program for the first two weeks, if possible, includes 20 minutes of an upright stationary bike, three hours of CPM usage (four hours if a bike is unavailable), passive range of motion exercises for 20 minutes and 20 repetitions of the isometric exercises twice each day. Passive range of motion maneuvers are important for decreasing postoperative stiffness and the development of adhesions while isometric exercises have the advantage of preventing muscular atrophy and hip flexor tendonitis.19,26 Additionally, Voight et al. found gluteal isometrics valuable for preventing iliopsoas (IP) spasms and decreasing anterior hip pain.10 The patient is also encouraged to lie prone for two hours per day to prevent hip flexion contractures.27,28

At each physical therapy visit in this stage, the patient begins on the upright stationary bike for 20 minutes followed by soft tissue mobilization (STM) for 20-30 minutes. Exercises at this stage include but are not limited to TA activation with bent knee fallouts, clamshell exercises, quadruped rocking, bridging progressions, and prone hip external and internal rotation (Figure 3). STM is designed to mobilize scar tissue and prevent soft tissue contractures with emphasis placed on the adductor musculature, quadratus lumborum (QL), tensor fascia lata (TFL) and iliopsoas (IP) (Figure 4). Additionally, the authors have found that resisted standing shoulder extension and alternating scapation, when combined with gluteal isometric activation is useful for improving core recruitmentand stability in a closed chain position (Figure 5). As Phase one progresses, the patient is weaned from crutches beginning at three weeks postoperatively, unless microfracture was performed in which case crutches are maintained until the six-week time point. Com-

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Figure 3. Exercises involved in Phase one of rehabilitation. Clams (A), knee fall outs with elastic support (B), bridging progressions (C-F), prone internal rotation (G), and prone external rotation (H).

Figure 4. Soft tissue mobilization strategies. STM is particularly useful in Phase one for the adductor musculature (A-B), iliopsoas (C), and quadratus lumborum (D).

mon pitfalls in this Phase include prolonged postoperative inflammation of soft tissue structures including the hip capsule and nearby musculature, hip flexor tendonitis, and lower back and sacroiliac joint pain. Lower back pain can often be managed with increased QL STM and avoidance of pelvic torsion. Hip flexor tendonitis often arises following premature active open chain hip flexor activation, and is commonly seen with toe touch or non-weightbearing gait patterns.26 Compensatory gait patterns are often responsible for postoperative lower back pain (LBP) and hip flexor tendonitis, making patient education on postoperative ambulation critical to a successful recovery.

ity to advance to closed chain exercises. Gait related restrictions that preclude progression include pain, a Trendelenberg, or compensated Trendelenberg gait, the presence of a trunk shift, and inadequate proximal control of the abdominal and lower extremity musculature. While specific criteria for advancement varies across institutions, there is general agreement that the patient should be in minimal pain, and ambulate without gait compensation, and display near normal range of motion with the contralateral leg before progressing to Phase two.17,20

While the authors do not employ formal criteria for advancement, the patient must demonstrate proper ambulation with a non-compensatory gait, and is encouraged not to push through pain in order to speed rehabilitation time. Further, the patient should demonstrate pain control, decreased edema, adequate strength and preparedness for ambulation without an assistive device. In addition, the patient must have adequate proximal control and the abil-

Phase Two: Return to Full Weight Bearing The goal of Phase two is to ensure the transition to full weight bearing progresses in a steady manner without the onset of deleterious compensatory strategies. (Table 2) The authors have found that the most difficult aspect of this phase is successfully weaning the patient off crutches. Aquatic therapy can assist in getting patients off crutches and can be started in week three, provided that the surgical incisions are completely healed. While not a requirement for the

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stages of this Phase include quadruped rocking and alternating shoulder flexion with hip extension, standing weight shifts in all directions, standing double leg knee bends (mini-squats), prone hip extension, and tall kneeling with an emphasis on core activation for all exercises (Figure 6). Additionally graded joint mobilization can begin with posterior and inferior glides beginning in week five (Figure 7). Importantly, STM for at least 20 minutes per session is continued through this phase in order to treat specific ROM limitations, prevent soft tissue irritation and tendonitis, and mobilize scar tissue around the portal sites.

Figure 5. Core strengthening exercises in Phase one include standing shoulder extension and alternating scapation with core and gluteal isometric activation.

rehabilitation protocol, aquatic therapy can be utilized as a supplement in the early stages of Phase two to improve mobilization and neuromuscular control as well as correct gait.29 Other exercises in the early

The second half of Phase two is designed to transition the patient to Phase three. It begins at week 6 with the introduction of both forward and backward elliptical training starting at low resistance for short durations. Additional exercises in the latter stage of this phase include single leg balancing with isometric abduction of the nonsurgical limb, eccentric psoas exercises, pole walks, internal oblique pulls, weightbearing hip internal and external rotation, functional walks with co-contraction of the core, gluteals, and gastrocnemius, as well as single leg stance with contralateral hip flexion and extension (Figure 8). One major pitfall of this phase is over-activation of the hip musculature as the patient gradually increases activity and hip range of motion. One specific example is with the development of hip flexor tendonitis,

Table 2. Phase Two Rehabilitation

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Figure 6. Exercises for the early stages of Phase two include rocking and alternate lower extremity extension and shoulder ďŹ&#x201A;exion in the quad position (A-B), core activation with half kneeling (C), and weightbearing isometric gluteus medius activation (D). Figure 7. Posterior and inferior joint mobilization can begin in Phase two of rehabilitation with inferior belt femoral glide (A), lateral to medial/inferior femoral glide (B), and medial to lateral femoral glide (C).

Figure 8. Later in Phase two, rehabilitation exercises begin to incorporate more functional movement about the hip. These include single leg balance of the surgical side with isometric abduction of non-surgical side (A), single leg stance with contralateral hip ďŹ&#x201A;exion and extension (B-C), functional walks with co-contraction of the core, gluteals, and gastrocnemius (D), and pole walks (E-F). The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 689

Table 3. Phase Three Rehabilitation

where weakness in the iliopsoas following arthroscopy results in hyperactivation and irritation of secondary hip flexors such as the rectus femoris, tensor fascia lata, and sartorius.15 These pitfalls can be minimized with STM at each therapy session as well as tactile and verbal cueing during exercises to promote proper muscle activation and technique. The patient can progress to Phase three when they are able to successfully demonstrate good proximal and distal control, avoid compensatory movements secondary to fatigue, and have minimal pain (pain 1-2 points higher than baseline on a 1-10 scale) throughout the entirety of the therapy session. Phase Three: Return to Pre-Injury Function Phase three is designed to return the patient to their pre-injury functional level with emphasis placed on dynamic exercises in all hip motion planes (Table 3). Lasting generally from weeks 8-16, it is more individualized than Phases one and two, and can last longer for patients with higher demands on their hip. In this phase, lunges, split squats, side steps, and retro-walks are instituted to complement the single leg balance, squat, and trunk rotation carried over from the end of Phase two. Joint mobilization and STM are continued on an as needed basis to manage isolated cases of pelvic malalignment, muscle over-activation and used to prevent tendonitis. Additional exercises at this stage include side planks with opposite lower extremity triplanar motion, resisted squats with Theracord rows, and core roll ups without flexor activation (Figure 9).

Figure 9. Exercises for Phase three include introducing rotational elements into hip motion (A-B), as well as Theracord assisted squats (C-D).

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The most common pitfalls during this phase result from rushing to return to activity. These pitfalls can manifest as decreased soft tissue mobility, recurrence of gait or running abnormalities, increased reports of pain, hip girdle tendonitis, or recurrence of pelvic malalignment. To prevent these setbacks, the therapist should watch for the presentation of increased femoral internal rotation, excessive pronation, hip drop/genu valgus, as well as an unequal arm swing as running is increased. If these deleterious motions are observed, running progression can be slowed to prevent the onset of compensatory and counterproductive running mechanics. The patient can progress to Phase four when they are able to complete all of the Phase three exercises without pain and can maintain good proximal and distal control with both running and other functional activities.

Figure 10. Triple Extension during running in Phase four.

Later in this phase, hip rotational exercises are introduced, while agility drills are deferred until Phase four. A treadmill walking program can begin at week 12, which can also be supplemented by light running using the Alter-G M320 anti-gravity treadmill (AlterG, Fremont CA). The Alter-G is a valuable addition to the rehabilitation regimen as both velocity and bodyweight percentage can be independently controlled (Figure 10; . This allows the patient to start running at reduced bodyweight earlier on in the rehabilitation timeline. This can function as an augment to the patientâ&#x20AC;&#x2122;s cardiovascular fitness, accelerate the patientâ&#x20AC;&#x2122;s neuromuscular reeducation following surgery, as well as improve the patientâ&#x20AC;&#x2122;s confidence while running.30,31 By only altering one variable at a time, either bodyweight or velocity, the therapist can slowly and safely advance the patient to return to run while being careful to watch for compensatory running patterns. In a pilot study evaluating return to run following Achilles tendon repair, Saxena et al. reported that the Alter-G allowed patients to progress to full weight bearing and running two weeks sooner than with the treadmill alone.31

Phase Four: Return to Running and Sport The goal of Phase four is to return the patient to full participation in their sport or recreational activity, focusing in particular on power, endurance, and agility (Table 4). While this phase is often tailored to be sport specific for each athlete, common exercises include box jumps, lateral cutting, scissor jumps, and single leg squats. Most institutions recommend that the patient be able to pass a series of functional return to sport tests where the patient is able to consistently display pain free running, jumping, lateral agility drills, and single leg squats.32 While the authors do not require a formal return to sport test to pass Phase four, a thorough return to running evaluation beginning at 16 weeks in an anti-gravity treadmill (AlterG; Fremont, CA) or 20 weeks on a standard treadmill. To progress to running, the patient must be able to demonstrate completely asymptomatic gait as well as the ability to complete all Phase three exercises without excessive fatigue. Advantages of the AlterG include the ability to identify and correct incorrect or compensatory running movements in a limited weight-bearing setting. For running evaluations, it is important for the therapist to witness triple extension with full body weight bearing running. Triple extension includes concurrent hip extension, knee extension, and ankle plantar flexion and is necessary for a powerful push off with running. Hip extension is particularly important to ensure adequate stride length and maintain the hip in a relatively neutral

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Table 4. Phase Four Rehabilitation

Table 5. Criteria to advance

position. Verbal and tactile cues should be employed by the therapist to ensure the patientâ&#x20AC;&#x2122;s running form is mechanically sound.33 Common deficits include Trendelenburg and compensated Trendelenburg with single leg squatting, and a valgus knee angle with lateral agility testing. DISCUSSION Rationale for Capsular Closure Historically, the treatment of many conditions of the hip required an open surgical procedure; however, over the last two decades open procedures have largely be replaced with arthroscopic procedures. These arthroscopic techniques, in general, have not placed emphasis on the treatment of the hip capsule. In the case of arthroscopic FAI treatment, most of the early descriptions detailed extended capsulot-

omies and even capsulectomies for adequate visualization of femoral-sided cam deformities.34 Only recently has more attention been paid to the capsule during arthroscopic treatment of hip conditions.21,35 The current protocol was developed using multiple expert opinions and clinical commentaries on hip arthroscopy rehabilitation as guides.14,15,19,26,27,36 Additionally, clinical practice guidelines have been published for non-arthritic hip pain by the Orthopedic Section of the American Physical Therapy Association.37 The present commentary outlines a rehabilitation protocol for a high volume hip arthroscopy center similar to what has been presented in the literature in terms of general rehabilitation time frame and Phased postoperative progression. Modifications to previously existing rehabilitation protocols include techniques designed to address the closed hip capsule following arthroscopic correction of FAI.

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Table 6. Pearls/Pitfalls of rehabilitation following hip arthroscopy for FAI

In arthroscopic treatment of hip conditions without capsular closure, iatrogenic instability caused by non-repaired capsulotomy or capsulectomy has become an increasing concern.38–41 Recently, there have been case reports of hip dislocation and subluxation following arthroscopic surgery for FAI without capsular repair.42–45 Capsular defects following hip arthroscopy have now been reported as one of the causes necessitating revision surgery for continued pain and/or instability.46 The clinical importance of capsular repair has been illustrated by Frank et al. as they demonstrated increased sports specific activity following hip arthroscopy for FAI with complete capsular repair compared to partial capsular

repair.21 Additionally, Wylie and colleagues have recently reported on 33 cases of symptomatic instability after hip arthroscopy without capsular repair and improvement in patient reported outcomes two years after revision hip arthroscopy in which the capsule was completely closed.22 These previous authors concluded that the contemporary treatment of FAI should include central compartment chondrolabral repair, surgical correction of osseous FAI, as well as complete capsular closure to optimize hip functional outcome and return to athletic activity. The clinical evidence stressing the importance of hip capsule integrity has also been corroborated with numerous biomechanical studies.47–50

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Table 7. Graded Joint Mobilization

Modiﬁcations to Rehabilitation Following Capsular Closure Following capsular closure, an emphasis is placed on protecting the anterior capsule and labrum, as strain on these structures with excessive hip flexion or hip extension past neutral the can impair healing.25 Graded joint mobilizations are an important addition to the rehabilitation protocol as it promotes safe hip range of motion while minimizing the risk for postoperative stiffness and contracture. If the capsule is left open, mobilization is contraindicated due to the risk of increased pain and the potential for iatrogenic instability manifesting as hip subluxation or dislocation.45 With capsular closure however, hip joint mobilizations are recommended to prevent excessive capsular stiffness, muscular guarding, and postoperative pain.28 Mobilizations begin with Grade I and progress through Grade III as tolerated. Mobilizations include inferior manual or belt assisted femoral glides, anterior to posterior femoral glides, lateral to medial femoral glides, medial to lateral femoral glides, and posterior to anterior femoral glides (Figure 7). Grade IV-V mobilizations are contraindicated to avoid excessive capsular laxity following arthroscopic closure (Table 7). As patients have varying degrees of preoperative capsular laxity, before undergoing a mobilization protocol, the therapist should discuss with the surgeon the quality of the capsular intra-operatively, as well as the state of repair. Repercussions following inadequate joint mobilization following capsular closure include gait deficit, decreased range of motion and a delayed return to sport. Graded joint mobilization is commonly used in the conservative management

of FAI with an intact capsule, and we have found that following capsular closure, mobilization techniques reduce postoperative joint hypomobility and improve rehabilitation progression.51 Return to Activity and Sport Return to sport is an important outcome measure for the young adult population undergoing hip arthroscopy for FAI and can be complicated by rehabilitation setbacks. To return to high intensity athletics, the patient must be able to demonstrate adequate high velocity and low velocity strength capable of explosive power generation.26 Currently, there is limited data regarding validated return to sport assessments specific to hip arthroscopy, however the Vail Hip Sports Test ™, a modification of the Vail Sport Test ™, has been used as a functional assessment for returning to sport following hip arthroscopy.18,52 Scored out of a maximum of 20 points, exercises include single leg squats for 3 minutes (6 points; 1 point every 30 seconds), lateral and diagonal bounding for 100 seconds each (5 points; 1 point every 20 seconds), and a forward box lunge for 2 minutes (4 points; 1 point every 30 seconds). The Tuck-Jump test is another functional assessment designed for ACL rehabilitation that has also been modified for the hip arthroscopy population.18 While useful to assess overall athletic competence, hip specific tests that take into account the unique rotational profile of the joint may be more advantageous to assess return to sport in the future. It has been the authors’ experience that patients are ready to begin return to sport activities once they successfully demonstrate treadmill running with correct form and minimal fatigability.

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CONCLUSION Appropriate postoperative rehabilitation is critical for overall patient success and optimal long-term outcome. As surgical techniques continue to evolve, rehabilitation protocols must adapt to accommodate changes in the surgical procedure and ultimately provide the safest and fastest recovery of function for the patient. One such surgical modification has been the incorporation of routine capsular closure as part of the treatment of FAI in the young, active patient. This clinical commentary details the authors current four-phase physical therapy program and the modifications made to incorporate the addition of capsular closure to the treatment of FAI. Following a step-wise approach including joint protection, progression to full weightbearing, addressing strength and function, and performing sport-specific rehabilitation will prepare an athlete for the safe and expedient return to sport following hip arthroscopy for the treatment of FAI. REFERENCES 1. Clohisy JC, Baca G, Beaule PE et al. Descriptive epidemiology of femoroacetabular impingement: a North American cohort of patients undergoing surgery. Am J Sports Med 2013;41:1348–1356. 2. Bozic KJ, Chan V, Valone FH et al. Trends in hip arthroscopy utilization in the United States. J Arthroplasty.2013;28:140-143. 3. Lynch TS, Terry MA, Bedi A et al. Hip arthroscopic surgery: patient evaluation, current indications, and outcomes. Am J Sports Med. 2014;41:1174–1189. 4. Zaltz I, Kelly BT, Larson CM et al. A. Surgical treatment of femoroacetabular impingement: what are the limits of hip arthroscopy? Arthrosc J. 2014:39;99-110. 5. Gupta A, Suarez-Ahedo C, Redmond JM et al. Best Practices During Hip Arthroscopy: Aggregate Recommendations of High-Volume Surgeons. Arthrosc J. 2015;31:1722-27. 6. Nwachukwu BU, Rebolledo BJ, McCormick F et al. Arthroscopic Versus Open Treatment of Femoroacetabular Impingement: A Systematic Review of Medium- to Long-Term Outcomes. Am J Sports Med 2014;44:1062-8.

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previous hip arthroscopic surgery. Am J Sports Med. 2016:44;39-45. Harris JD, Slikker W, Gupta AK et al. Routine complete capsular closure during hip arthroscopy. Arthrosc Tech. 2013;2:89-94. Ayeni OR, Bedi A, Lorich DG et al. Femoral neck fracture after arthroscopic management of femoroacetabular impingement: a case report. J Bone Joint Surg, Am. 2011;93:e47. Domb BG, Sgroi TA, VanDevender JC et al. Physical therapy protocol after hip arthroscopy: clinical guidelines supported by 2-year outcomes. Sports Health 2016;8:347-354. Malloy P, Malloy M, Draovitch P. Guidelines and pitfalls for the rehabilitation following hip arthroscopy. Curr Rev Musculoskelet Med. 2013;6:235-241. Stalzer S, Wahoff M, Scanlan M. Rehabilitation following hip arthroscopy. Clin Sports Med. 2006;25:337-357. Cheatham SW, Kolber MJ. Rehabilitation after hip arthroscopy and labral repair in a high school football athlete. Int J Sports Phys Ther. 2012;7:173-184. Villalta EM, Peiris CL. Early aquatic physical therapy improves function and does not increase risk of wound-related adverse events for adults after orthopedic surgery: a systematic review and metaanalysis. Arch Phys Med Rehabil. 2013;94:138-148. Grabowski AM. Metabolic and biomechanical effects of velocity and weight support using a lower-body positive pressure device during walking. Arch Phys Med Rehabil. 2010:91;951-957. Saxena A, Granot A. Use of an anti-gravity treadmill in the rehabilitation of the operated achilles tendon: a pilot study. J Foot Ankle Surg. 2011:50;558-561. Domb BG, Stake CE, Finch NA. et al. Return to sport after hip arthroscopy: aggregate recommendations from high-volume hip arthroscopy centers. Orthopedics 2014;37:e902-905. Agresta C, Brown A. Gait retraining for injured and healthy runners using augmented feedback: a systematic literature review. J Orthop. Sports Phys Ther. 2015;45:576-84. Larson CM, Guanche CA, Kelly BT et al. Advanced techniques in hip arthroscopy. Instr. Course Lect. 2009:58;432-436. Bedi A, Galano G, Walsh C et al. Capsular management during hip arthroscopy: from femoroacetabular impingement to instability. Arthrosc J. 2011;27:1720-1731. Malloy P, Gray, K, Wolff AB. Rehabilitation after hip arthroscopy: a movement control–based perspective. Clin Sports Med. 2016:35;503-521. Enseki K, Harris-Hayes M, White DM et al. Nonarthritic hip joint pain. J Orthop Sports Phys Ther. 2014;44:A1-32.

38. McCormick F, Slikker W, Harris JD et al. Evidence of capsular defect following hip arthroscopy. Knee Surg Sports Traumatol Arthrosc. 2014;22:902-905. 39. Cvetanovich GL, Harris JD, Erickson BJ et al. Revision hip arthroscopy: a systematic review of diagnoses, operative ﬁndings, and outcomes. Arthroscopy. 2015:31;1382-1390. 40. Weber AE. Harris HD, Nho SJ. Complications in hip arthroscopy: a systematic review and strategies for prevention. Sports Med Arthrosc Rev. 2015;23:187-193. 41. Bayne CO, Stanley R, Simon P et al. Effect of capsulotomy on hip stability-a consideration during hip arthroscopy. Am J Orthop. 2014;43:160-165. 42. Matsuda DK. Acute iatrogenic dislocation following hip impingement arthroscopic surgery. Arthroscopy. 2009;25:400-404. 43. Ranawat AS. McClincy M, Sekiya JK. Anterior dislocation of the hip after arthroscopy in a patient with capsular laxity of the hip. A case report. J Bone Joint Surg Am. 2009;91:192-197. 44. Benali Y, Katthagen BD. Hip subluxation as a complication of arthroscopic debridement. Arthroscopy. 2009;25:405–407. 45. Duplantier NL, McCulloch P, Nho SJ et al. Hip dislocation or subluxation after hip arthroscopy: a systematic review. Arthroscopy 2016;32:1428-1434. 46. Philippon MJ, Stubbs AJ, Schenker ML et al. Arthroscopic management of femoroacetabular impingement osteoplasty technique and literature review. Am J Sports Med. 2007;35:1571-1580. 47. Abrams GD, Hart MA, Takami K et al. Biomechanical evaluation of capsulotomy, capsulectomy, and capsular repair on hip rotation. Arthroscopy. 2015;31:1511-1517. 48. Khair MM, Grzybowski JS, Kuhns BD et al. The effects of capsulotomy and capsular repair on the biomechanics of hip distraction. Arthroscopy. 2017;33:559-565. 49. Myers CA, Register BC, Lertwanich P et al. Role of the acetabular labrum and the iliofemoral ligament in hip stability an in vitro biplane ﬂuoroscopy study. Am J Sports Med. 2011;39:85S–91S. 50. Ito H, Song Y, Lindsey DP et al. The proximal hip joint capsule and the zona orbicularis contribute to hip joint stability in distraction. J Orthop Res. 2009;27:989–995. 51. Loudon JK, Reiman MP. Conservative management of femoroacetabular impingement (FAI) in the long distance runner. Phys Ther Sport. 2014;15:82-90. 52. Pierce CM, Laprade RF. Wahoff M et al. Ice hockey goaltender rehabilitation, including on-ice progression, after arthroscopic hip surgery for femoroacetabular impingement. J Orthop Sports Phys Ther. 2013;43:129-141.

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IJSPT

CLINICAL COMMENTARY

ENERGY SYSTEM DEVELOPMENT AND LOAD MANAGEMENT THROUGH THE REHABILITATION AND RETURN TO PLAY PROCESS Scot Morrison, PT, DPT, OCS, CSCS1,2 Patrick Ward, MS, CSCS, LMT3 Gregory R duManoir, PhD4

ABSTRACT Return-to-play from injury is a complex process involving many factors including the balancing of tissue healing rates with the development of biomotor abilities. This process requires interprofessional cooperation to ensure success. An often-overlooked aspect of return-to-play is the development and maintenance of sports specific conditioning while monitoring training load to ensure that the athleteâ&#x20AC;&#x2122;s training stimulus over the rehabilitation period is appropriate to facilitate a successful return to play. The purpose of this clinical commentary is to address the role of energy systems training as part of the return-to-play process. Additionally the aim is to provide practitioners with an overview of practical sports conditioning training methods and monitoring strategies to allow them to direct and quantify the return-to-play process. Level of Evidence: 5 Key words: Acute/chronic workload, energy system development, return-to-play, sports rehabilitation, strength and conditioning, training load

Professional Referee Organization, New York, NY, USA Black Diamond Physical Therapy, Portland OR, USA 3 Seattle Seahawks, Seattle, WA USA 4 The University of British Columbia Okanagan, Kelowna, British Columbia, Canada 2

CORRESPONDING AUTHOR Scot Morrison Black Diamond Physical Therapy 4224 NE Halsey, Suite 340 Portland, OR 97213 Phone: (503) 288-4643 Fax: (503) 208-7016 E-mail: scotmorrsn@gmail.com

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INTRODUCTION Injury is an unfortunate outcome for some of those involved in sports. An estimated seven million Americans receive medical attention for sports related injuries each year.1 Once an athlete sustains an injury it is important that they progress through a focused rehabilitation program to allow them to return to their previous level of competition. This process has been termed the return to sport continuum and is separated further into three distinct phases (Figure 1).2 The approach taken during this transition from injury to full participation is critical in preparing the athlete for competition at a high level while also reducing the risk of re-injury.3 Unfortunately, the return-to-play (RTP) approach is not always handled in a systematic way. This can lead to a delay in the athlete’s RTP, leaving the athlete underprepared for the demands of sport with increased risk of re-injury or decreased performance upon return.4-6 The RTP process is a multi-disciplinary venture including sports physicians, physical therapists, athletic trainers, strength and conditioning coaches, and sport coaches. Collectively, these professionals as well as the athletes themselves all play a role in the decision making process.2,7 Communication amongst these groups is paramount in optimizing athletic care and ensuring the best possible outcomes.2,5 However, the criteria used to gauge progress are often vague and lack standardization.8 As a result communication during the RTP process can be compromised. In some instances a standard of practice is used that focuses on local tissue adaptations based on healing time frames.8,9 While a time

Figure 1. Return to Play Continuum .

based approach does improve communication, Herrington and colleagues make a strong argument that this is insufficient and a task-based approach that focuses on clearly defined performance goals is the best approach for RTP.9 Thus it is important that the language that is used in describing the RTP process is familiar to everyone involved, is measurable, and reflects best practices. As an athlete recovers, the emphasis progresses from protecting the injured tissue, to guiding the healing process, and finally into restoring the capacity of strength and energy systems. However, this restoration of strength and capacity may not be fully realized during the athlete’s rehabilitation.10 In the best-case scenario, an athlete will be “medically cleared” to play using RTP algorithms (e.g., muscle strength, joint range of motion, a series of performance tests, etc.), which have been shown to be of value in making this decision.11 However, even these approaches typically give very little information regarding how much training the athlete has performed or whether their fitness is sufficient to tolerate competition at a high level.12,13 Better documentation of the training process during RTP allows for the quantification of workload, providing direction to the program and enhancing communication. This need for an assessment of the workload performed in the RTP process and its potential link to re-injury was recently listed as a key focus area by the 2016 Consensus Statement on Return to Sport from the First World Congress in Sports Physical Therapy, Bern.2 To address some of these concerns several authors have proposed RTP frameworks, which can aide in discussions amongst the relevant practitioners.4-7 For example, the updated Strategic Assessment of Risk and Risk Tolerance (StARRT) framework breaks the decision making process down into three parts: 1) the assessment of health risk, 2) the assessment of activity risk, and 3) the assessment of risk tolerance.4,6 Approaches like this establish a system that monitors the entire process while assigning each professional a role in clearing the athlete for RTP. None of the approaches referred to above have explicitly described an assessment of the athletes acute or chronic workload. However the authors of this paper find that the StaRRT framework suits this process well and propose that the assessment

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of workload should be included in each of the three categories laid out in the StaRRT framework.2,4,6 This will allow for a smooth transition from injury to medical clearance for RTP and on to return to performance while ensuring that the athlete has performed an appropriate amount of chronic loading to tolerate these progressions.5,6,12,14

vided to working muscle through the interactions of the three metabolic pathways:16,17

It is typical that strength and conditioning is performed by “healthy” athletes.5 However, injured athletes also benefit from a training program that prevents detraining and helps to re-develop any biomotor adaptations lost during the initial rehabilitation process. This thought process has led various authors to propose a comprehensive approach to strength and conditioning within the RTP paradigm.5,14,15 The role of strength training within rehabilitation has been discussed elsewhere;14,15 but very little has been described regarding re-conditioning, or energy systems development during the RTP process. All of the bodies’ energy systems play a supporting role at the onset of activity thus any form of training can be used to promote energy system development. However, specific acute training variables such as mode of exercise, duration, intensity, and rest interval dictate the type of adaptations elicited.

The intensity and duration of the exercise bout determines which of the energy systems has the greatest contribution to energy supply. Activities such as Olympic Weightlifing and the 100m sprint utilize ATP at a very high rate and therefore rely on the Adenosine triphosphate-Phosphocreatine (ATPPCr) pathways to rephosphorylate ADP to ATP. This pathway relies on the single-step creatine kinase reaction to rephosphorylate ADP and thus has a high rate of energy supply. However, there is a finite supply of intramuscular ATP and PCr stores resulting in a very limited capacity of this system.

Training and rehabilitation are complex processes that require sufficient monitoring to ensure that performance objectives are being met. The goal of the training process is to ensure that the training is progressed at the optimal pace and that the athlete is not under or over exposed to training loads.12 As such, all individuals involved in RTP decision-making process should have an understanding of the athlete’s current level of fitness and its relation to their clearance for RTP. The purpose of this clinical commentary is to address the role of energy systems training as part of the RTP process. Thus, an overview of the three metabolic pathways followed by energy system training concepts will be provided. Finally, practically applicable methods are given for monitoring the training process providing practitioners with tools to quantify the training result and direct the RTP process for the athlete. ENERGY SYSTEM PHYSIOLOGY Exercise places both metabolic and neuromuscular demands on the body16 and in order to meet these demands adenosine triphosphate (ATP) must be pro-

1. The phosphagen system 2. The glycolytic system 3. The oxidative system

Intermediate duration activities such as 400m sprinting must rely on higher capacity energy systems to rephosphorylate ATP. This multistep pathway utilizes glucose/glycogen to provide ATP for continued muscular contraction and will so in the absence of oxygen. However, fatiguing by-products are produced, which limit the duration of this energy system. Longer duration activites such as marathon running rely on oxidative phosphorylation pathways (aerobic glycolysis/β-oxidation, Kreb’s Cycle, electron transport chain) to sustain ATP production over the extended exercise duration. However, these pathways require the integrated delivery and utilization of oxygen and thus have substantially lower rates of ATP supply. Although not mutually exclusive, the proportional reliance on these three energy systems depends on both the duration and intensity of exercise (Table 1). Table 1 may serve to provide a general understanding of the metabolic demands of various sport activities. For example, the metabolic requirements of sports such as Olympic weightlifting, 100m sprinting, or Marathon running, are relatively easy to classify since the intensity and specific durations of these sports biases them towards the ends of the spectrum. The demands of many team sports, however, present a greater challenge when classifying them based on metabolic or energy systems

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Table 1. Energy System Demands

Duration

Intensity

Primary Energy System

0 – 6 seconds

Maximal

Phosphagen

6 – 30 seconds

Near-Maximal

Phosphagen & Anaerobic Glycolysis

High

Anaerobic Glycolysis

Moderate

Anaerobic & Aerobic Glycolysis

Low

Oxidative

30 – 120 seconds 2 – 3 minutes > 3 minutes

demand. A majority of team sports are comprised of maximal or near maximal sprints followed by brief periods of recovery18 and therefore do not utilize the extremes of the energy system as in the previous examples. Descriptions of metabolic pathways, such as Table 1, make it appear as though these pathways are mutually exclusive with distinct ‘on-off’ switches as the duration increases and the intensity of work decreases. It should be noted that each of the systems works to supply energy in concert, however their relative contributions change depending on the task.19 As such, team sport athletes require fitness levels sufficient to support metabolic requirements specific to their sport and position that spans the three main metabolic pathways. The “on and off”, intermittent nature of team sport has been termed repeated sprint ability (RSA).20,21 RSA reflects a key characteristic of team sport, whereby a large portion of competition is spent performing lower intensity activities with brief bouts of maximal to near maximal efforts interspersed throughout the time played.20 Thus, it is important that athletes participating in activities with high RSA demands possess the ability to repeat maximal or near maximal efforts with limited fatigue.17,20 As fatigue accumulates over the course of a competitive match the ability to repeatedly sprint at a maximal speed becomes compromised. Multiple mechanisms of fatigue have been explored within the construct of RSA.17,20,22,23 In addition, many acute training and competition variables have been studied as well. For example, decreased RSA due to fatigue may be attributed to duration, work-to-rest interval,24 type of recovery (active vs. passive),17,25 oxygen uptake kinetics26 and

changes in muscle metabolites.27 While the physiological underpinnings of RSA continue to be explored it is important to note that a variety of training methods have been proposed to address these limitations and improve RSA.17,28 For example, high-intensity interval training (HIIT) methods have been shown to improve VO2max,29 VO2 kinetics,30 mitochondrial biogenesis,31,32 and sports performance.33,34 Energy system training terminology is not always consistent, which may result in confusion in the application of research principals to the actual training methods used.35 To alleviate this, Chamari and Padulo recommend doing away with physiological descriptions of training and instead propose three basic classifications for short duration – repeated effort exercise. This classification system can help practitioners develop training sessions without being confused or hindered by exercise physiology terminology. The classification system proposed by Chamari and Padulo are based on exercise duration and can be seen in Table 2. Note that longer duration exercise is not addressed within the classification structure of Chamari and Padulo; however, it should not be overlooked. While excessive amounts of long duration, endurance exercise may have deleterious effects on the explosive capabilities of power athletes36 a moderate amount of this type of conditioning has been found to increase performance in team sport competition as a major portion of team sport is spent performing low-intensity activities in-between high intensity efforts.37 During these periods, the aerobic energy system plays a central role in recovery between

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Table 2. Energy System Training Distribution

Classification

Intensity

Duration

Programming Example

Explosive Efforts

Maximal

Up to 6 s

3-5 s work : 60-120 s Rest

High Intensity Efforts

Maximal

6 s to 1 min

30 s work : 30 s Rest

Endurance Efforts

Maximal

> 1 min

3 min work : 3 min Rest

intense bouts of exercise and assists in repeating those high intensity efforts with less performance decrement.38-40 While the aerobic system may contribute 10% or less energy to a single sprint, with repeated sprints its contribution can rise to as much as 49%.38 For these reasons, a well-rounded training program should be designed to ensure that the full spectrum of the athleteâ&#x20AC;&#x2122;s fitness is addressed within the RTP program. TRAINING METHODS & PROGRAMMING A determination of the individual needs of the athlete is important to ensuring that the testing done during the RTP process accurately assesses the demands that will be placed on them. This is done with a needs analysis of the sport and should be specific to the position the athlete plays as well as the level of competition they participate in (e.g., collegiate, amateur, professional). An in depth discussion of this processes is beyond the scope of this paper and the reader is referred to the following papers for an exploration of this topic in full.41,42 Numerous factors must be considered when deciding what to

emphasize within training, such as the time since the initial injury, what has been done to date, and the time frame available before RTP. Communication between professionals involved in the RTP decision at this point will help to ensure that the RTP process is optimized. The manipulation of acute training variables dictates the response the athlete has to the training program. Buchheit and Laursen23 recognize nine variables that can be manipulated within an energy system training session (Table 3). The manipulation of training variables can be used to change the focus of the session or to aid in selecting modalities (e.g., cycling, rowing) that may be appropriate during phases of the RTP process when activities like running are contra-indicated. These variables offer the practitioner a number of options for creating a training program. Table 4 provides guidelines for the manipulation of these variables based on the objectives of the training session (e.g., short duration or long duration) and training modalities that might be appropriate for the individual at that time. The

Table 3. Energy System Training Variables

Segment of Training Session

Work Interval

Recovery

Series: A group of sets performed in succession before a longer rest break is taken.

Variable Manipulated Intensity Duration Modality (e.g., Run, Bike, Rowing, etc.) Duration Intensity Number of Series Series Duration Time Between Series Between-Series Recovery Intensity

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Table 4. Energy System Training Parameters

Work Interval Intensity

Duration

Recovery Intensity

Series

Duration

Number of Series

Adaptation

Weekly Frequency

Timeframe

Short Duration â&#x20AC;&#x201C; Repeated Efforts Explosive Effort

Maximal

<6 s

Passive

30 to 120 s

2 to 6

2 to 3

2 to 3 weeks

HighIntensity Effort

Maximal

15 to 30 s

RPE <2

30 to 120 s

4 to 10

2 to 3

2 to 3 weeks

2-3 min

RPE <2

2-3 min

6 to 10

2 to 3

2 to 3 weeks

3 to 5

2 weeks to 3 months

2 to 3

2+ weeks

Endurance RPE Effort 8 to 9

Long Duration â&#x20AC;&#x201C; Endurance Extensive

Zone 1

20-60 min

Intensive

Zone 2 Zone 3

6-8 min 4-6 min

Continuous Low Zone 1

clinician is advised to utilize this framework when designing programs. When designing training programs, the practitioner is required to determine the appropriate training intensity for a given training session. Prescription of intensity has been based on of a number of physiological measures in the literature including maximal heart rate, VO2max, aerobic and anaerobic threshold, and critical power.43 Traditionally a percentage of these values are utilized to prescribe an internal (%VO2max, % threshold) or external (% critical power) intensity in either a continuous or interval-based prescription. With the popularity of interval-based prescriptions23 (see Training Methods & Programming) it is imperative that the practitioner not only considers the intensity of the exercise bout, but also the intensity of inter-interval recovery periods and inter-series recovery. Technology, such as GPS monitors, heart rate monitors, and power meters, may assist practitioners in directly quantifying training intensity, however, expense may

2-4 min

3 to 6

limit their applicability in certain situations. The use of perceived intensity and duration make for an affordable solution that is easy to implement in any setting. The rating of perceived exertion (RPE) is one of the most recognized measures for monitoring training intensity during a workout.44 Using this method clinicians may then prescribe exercise intensities within a binary or 3 zone model as shown in Table 5.44-47 The binary zone model differentiates low vs. high-intensity with respect to the first blood lactate threshold (2.5 mmol/L) or 72% HRmax, whereas the 3 zone model includes intensities below 2.5 mmol/L blood lactate or 55-82% HRmax as zone 1, 2.5 mmol/L to 4.0 mmol/L blood lactate or 82-87% HRmax as zone 2, and >4.0 mmol/L blood lactate and >87% HRmaz as zone 3.47 The RPE scale most often used is the modified CR10 scale seen in Table 7.45 These parameters offer a flexible model for developing both anaerobic and aerobic qualities in the injured athlete.

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The classification system established in Table 2 gives the practitioner a framework that can be used for designing and progressing the individual rehabilitation program. Key training sessions throughout the week would be designed to improve the predominant energy system demands required by the sport and position the athlete competes in. For example, the sport of American football is an intermittent sport comprised of repeated, high-intensity efforts of approximately 4 – 7s, followed by 15 – 80s of recovery, which would be considered “Explosive Effort” per Table 2.48-50 As such, the goal of the programming would be to maximize the various energy systems ability to tolerate these demands. Table 6 provides an example program demonstrating application of the various principals to a specific case. Training modalities should progress as the athlete moves through the

RTP program. For example, the athlete may begin performing bike workouts, if running is initially contraindicated due to their injury, and then progress to linear running, change of direction running, and finally to open environment skill work, where the athlete must sprint at a high level whilst making decisions and changing directions. Similarly, exercise intensity would progress from lower intensity work to explosive efforts of straight ahead running, to change of direction work performed under similar work-torest ratios specific to the game. Finally, the athlete’s program would be progressed from longer rest intervals to rest intervals that are more sport specific. MONITORING THE PROCESS Monitoring the training process provides practitioners with the ability to quantify the volume and

Table 5. Sustained Effort Training Intensity Zones

Table 6. Sample Basketball Sports Speciﬁc Training Progression after Injury to Lower Extremity (LE)

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intensity of training an athlete has performed. In doing so, the practitioner can methodically manipulate training variables and increase or decrease the amount of training stress on a given day to allow for consistent improvement without exposing the athlete to loads that they are not prepared to tolerate. Several methods have been proposed to assist practitioners in understanding the dose-response relationship of training.51-54 In order to provide practitioners with low cost methods, this commentary will focus mainly on the use of Session Rating of Perceived Exertion (sRPE). More objective methods of training monitoring exist (e.g., HR, Global Position System (GPS) tracking, etc.); however these methods can be costly and require specific expertise to handle the data is required. QUANTIFYING TRAINING SESSIONS The dose-response relationship describes the interaction between what the athlete did in training and how they responded to it.51 A distinction between external (what the athlete did) and internal (how they responded) training load factors allows for contextualization of the applied training loads. External training load represents the work performed by the athlete in a given training session.53 It can be quantified in several ways, such as the distance covered, amount of high speed running performed, weight lifted, or total training volume. The rehabilitation specialist or strength coach plans the expected external training load for a given day; as they select the training modality, exercise intensity, volume of work, and the work to rest ratio. Various methods of external load quantification have been

Name Date

Athlete A 1/4/16

Training Modality Method

Hill Sprint

Intensity Maximal

Series 4

Training Duration

51 min

explored in the literature.53 The use of integrated microtechnology (GPS, accelerometers, and gyroscopes) has become one of the most popular methods of external load quantification in team sport53,55-57 as these systems allow for the quantification of running, collisions, accelerations, decelerations, and change of direction measures during practice or sports training.56,58,59 Unfortunately, the expense of these systems may make their use impractical for practitioners in private practice or those working with teams or smaller universities with limited budgets. In these cases charting the details of the athletes training session is a simple alternative to track external training load as seen in Figure 2. The athleteâ&#x20AC;&#x2122;s response to the planned training session is termed internal training load.51,53 A well documented method of internal load quantification is heart rate response.51,53,54,60 The relationship between heart rate response and exercise intensity during a training session has been explored using various training impulse (TRIMP) models, which assign a weighting factor to an arbitrary number of HR zones.51,53,54,60 The HR zones are then summated to create a training load score for a given training session. While this approach is easy to apply the financial burden of purchasing equipment is still present60 and the kinetic response of HR adjustment makes them impractical for interval based training. Thus the use of internal load quantification via the session Rating of Perceived Exertion (sRPE) method is likely the most clinically applicable method available.45,51,53,54,60 Originally developed by Foster and colleagues,45,51 the RPE method allows the quantification of training

Explosive Efforts Rest Between Series 5 min

Sets 15

Work 5 sec

Rest 30 sec

Figure 2. Example Training Document. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 704

load without the use of heart rate technology. Foster and colleagues found a consistent relationship between TRIMP scores and the individual’s RPE during aerobic exercise45 and others have found similar relationships with RPE and %HRpeak and blood lactate.61 The athlete is asked to rate the intensity of the training session utilizing the descriptors of the CR10 Scale (Table 7).45,62 An arbitrary training load based on session RPE can be calculated by multiplying the corresponding number for the descriptor on the CR10 Scale by the duration of training, in minutes.62,63 These values can be stored in any spreadsheet software file for the individual athlete and changes in training load can be charted over time (Figure 3). It has been recommended that the athlete be asked to provide their session RPE approximately 20-30min following the training session.63,64 The timing of the response is thought to improve the athlete’s ability to reflect on the session as a whole versus their perception of the last activity performed in the training session, which could be skewed by their overall fatigue at that time.45,64 This limitation may present a problem for practitioners, as keeping the athlete in the facility for 20-30min following their training session is often not practical. More recently, the importance of measurement timing has been called into question. Uchida and colleagues65 collected sRPE on boxers at both 10 and 30 min post training following three standardized training ses-

Table 7. CR-10 Rating of Perceived Exertion

Rating 0 1 2 3 4 5 6 7 8 9 10

Descriptor Rest Very Easy Easy Moderate Somewhat Hard Hard Very Hard Very, Very Hard Almost Maximal Maximal

sions (easy, moderate, and hard). They found that RPE had little variation between the 10 and 30min measurements across all three training sessions. Kraft and colleagues66 observed a similar response when quantifying sRPE for resistance training sessions taken at 15 vs. 30min. These findings indicate that practitioners may be able to practically apply the sRPE method in settings where it is not practical to wait 20-30min following training.

Figure 3. Change in Training Load Over Time, for hypothetical “Athlete A”. Note: Values are calculated using RPE x Time. The International Journal of Sports Physical Therapy | Volume 12, Number 4 | August 2017 | Page 705

Initially developed for aerobic exercise, session RPE has also been found to be valuable for quantifying other training activities. For example, session RPE has been used as a marker for training load during intermittent team sport activities and resistance training, providing practitioners with an affordable and easy to administer method of quantifying the internal training response.60,66-73 Finally, in addition to quantifying the session as a whole, RPE can be used to quantify bouts of work during energy system training.28 If the practitioner lacks a HR monitor, a prescribed RPE could be presented to the athlete for their work intervals. For example, performing high intensity efforts at an RPE of 8-9 for 30 s of work followed by 30 s of recovery at an RPE of 2. MODELING THE TRAINING PROGRAM Daily training loads can be evaluated over time to quantify how much training the athlete has performed during their RTP process and to ensure that they have trained enough in order to withstand the loads of practice and competition. As a method of understanding periodization in sport, RPE has often been reported in the literature in absolute terms, reflecting cumulative loads over one to two week blocks of training, or percent changes from one week to the next.67,68,73-75 While this method of reporting provides a useful look at how an athlete’s program is progressing it tells little about the cumulative effect that training has had on the body. Utilizing RPE, a recent method proposed by Hulin and colleagues seeks to take into account both acute and chronic load during the training process as a way of quantifying changes in fitness and fatigue.76 This method, termed the acute:chronic ratio, allows practitioners to determine how the athlete is tolerating the training processes over time, while slowly progressing back to pre-injury fitness levels.12 The acute load represents the most recent weekly training load for the athlete while the chronic load is the four-week average of work the athlete has performed.76 The acute load is divided by the chronic load to produce a single number, which describes the athlete’s current status.77 A ratio greater than 1 would indicate that the acute load for that week exceeded what the athlete had previously been exposed to.76

Using this method, it was observed that rugby athletes who exceed an acute:chronic workload of 2.11 were at a 3.4 times greater risk of injury.77 The metric is easy to compute and a visualization of the training response during the RTP program can serve as a discussion point amongst medical, rehabilitation, strength and conditioning coaches, and coaches. Table 8 shows an example of using the acute:chronic training load to plan training. Periods of intense training are helpful in building an athlete’s fitness, however, excessive periods of intense loading expose the athlete to increased risk of injury or, for the rehabilitating athlete, re-injury.45,63,78,79 The acute:chronic training load is useful for identifying periods of excessive loading, as can be seen by the conditional formatting in Figure 4. The acute:chronic ratio is flagged “red” any time the acute:chronic training load exceeds 2 and “yellow” any time the acute:chronic load exceeds 1. For example, in the first case there is an acute:chronic load of 2.10. The acute load in that given session was 6856, which is 2.10x greater than the average of the previous four sessions (3269.3). This data can be represented graphically as well. The bottom of Figure 4 shows a chart representing the athlete’s acute load (red), chronic load (green), and acute:chronic ratio (grey shading). Additionally, the threshold lines indicating an acute:chronic load of 1 and 2 are also represented. In the example the practitioner noticed the high training loads for three consecutive weeks (weeks 5-7) before planning a few weeks of lower training (weeks 8-10) to allow for recovery and a dissipation of accrued fatigue. CONCLUSION RTP from an injury is a challenging process for everyone involved. The numerous factors inherent to decision making within the RTP process have been addressed in the recently updated StARRT framework.4,6 The framework is defined by a comparison of the individuals risk tolerance to the risk involved in their RTP. This clinical commentary has attempted to demonstrate how this can be used when assessing and programming energy system development in the injured athlete. An assessment based on the demands of the sport and the acute:chronic training load can be utilized to monitor the gap between these two domains. The implementation of energy system development principals allows

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Figure 4. Acute: Chronic Training Load for hypothetical “Athlete A”. Note: Values are calculated using RPE x Minutes

the clinician to address the athlete’s readiness and tolerance to the increasing demands they will face. By approaching the return to sport question in this manner the RTP team is able to ensure that relevant risks specific to capacity and workload are discussed and addressed during the rehabilitations and RTP process. This allows the athlete to RTP with a capacity that is able to handle the levels of stress that their spor requires.

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