Biologically based design of an actuator systemfor a knee–ankle–foot orthosis

Mechanism and Machine Theory 44 (2009) 860–872

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Biologically based design of an actuator system for a knee–ankle–foot orthosis A. Cullell, J.C. Moreno *, E. Rocon, A. Forner-Cordero, J.L. Pons Instituto de Automática Industrial, Consejo Superior de Investigaciones Científicas – CSIC, Carretera Campo Real Km 0,200, La Poveda – Arganda del Rey, Madrid, Spain

a r t i c l e

i n f o

Article history: Received 10 October 2007 Received in revised form 1 April 2008 Accepted 7 April 2008 Available online 19 May 2008

Keywords: Actuator Biomechanics Human gait Knee–ankle–foot orthosis Rehabilitation Wearable robot

a b s t r a c t Knee–ankle–foot orthosis are systems used to restore human gait, providing stability during stance phase. A concept of actuator for knee–ankle–foot orthosis was developed based on biomechanical data. The actuator is conceived to provide mechanical means to reproduce the normal kinematics during human gait at joint level. Behaviour of the joints of the lower limb was approximated by elastic means and an actuator for each joint was designed and constructed. The rationale of the design process is presented, considering the functional aspects and aiming at a lightweight solution with low power demand. Tests performed with one patient suffering from post-poliomyelitis syndrome are presented and evidence of functional compensation during stance and swing phases with the proposed solution is given. Ó 2008 Elsevier Ltd. All rights reserved.

1. Introduction Muscular weakness of the lower limb can be the result of different diseases, such as peripheral neurological diseases (poliomyelitis and post-polio syndrome, spina bifida, poly neuropathy), muscular diseases (duchenne muscular dystrophy and Becker’s muscular dystrophy, myasthenia gravis), and central neurological diseases (multiple sclerosis, cerebral palsy, brain injury, cerebrovascular accident, spinal cord injury). The percentage of people suffering from these diseases can oscillate between 0.05% and 1% of the total European population [1,2]. The result of insufficient muscle-strength or muscle-control can be either a recurvated knee while standing or anterior knee instability during stance resulting in risk of falling.

Knee–ankle–foot orthoses (KAFOs) are prescribed as a partial solution for these disorders while providing stability and keeping joints at their functional position. Conventional systems provide stability during walking by maintaining the knee in a fixed position, but this results in unnatural gait patterns. While the user walks with a locked knee, there is no possibility to flex the knee during pre-swing and swing phases. As a result, the user has to swing the body to the side of the non-affected leg to clear the swinging leg. Besides, having the knee locked during the complete gait cycle requires excessive energy consumption. There is also absence of loading response during knee flexion. Different systems that release the knee during swing phase, allowing the free oscillation of the leg due to inertia [3–5], have been proposed. These systems avoid swinging of the body during gait, but have not yet probed advantages in terms of energy consumption. Another type of device is the

* Corresponding author.

E-mail address: moreno@iai.csic.es (J.C. Moreno).

0094-114X/$ - see front matter Ó 2008 Elsevier Ltd. All rights reserved. doi:10.1016/j.mechmachtheory.2008.04.001

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hybrid KAFO, which combines functional electrical stimulation (FES) with mechanical systems [6,7]. Electrical stimulation is applied to the muscles of the leg to improve its trajectory during swing phase, and an orthosis is responsible to provide the necessary torque during stance by means of brakes or clutches, which lock the knee when required. Discomfort related to stimulation is the major problem of this technique. The system developed by Kameyama and Suga [8] controls joint movements with a brake driven by a servomotor, which adjusts the strength of the knee fixation to the requirements of the gait cycle.

Other passive systems can modulate the impedance of the joint. One example is the ankle–foot orthosis by Blaya and Herr [9]. In this system the impedance of the orthotic joint is modulated throughout the walking cycle by changing the stiffness of a controllable actuator.

The prosthesis developed by Carlson et al. [10] modulated the impedance of the joint as well, but in this case it is the damping of the joint that can be regulated by changing the properties of the magnetorheological fluid of the actuator. Other systems use active actuators (pneumatic [11,12] or electric [13,14]) to control joint movements, not only controlling its position, but providing the needed action when necessary. One important restriction of these systems is the necessity of an external power source, which limits the use of the system to the laboratory.

Analyzing the existing orthotic systems for the lower limb it becomes clear that the main problems still present in state of the art of such systems are unnatural gait patterns, energy consumption, and the need of an external power source are main drawbacks. Traditional and commercially available knee–ankle–foot orthoses, consist of three lateral bars connected by two hinges (knee and ankle), fixated at the anatomical parts (see Fig. 1).

A KAFO for functional compensation of lower limb joint disorders was developed. The system uses a storing-releasing energy concept to apply functional compensation on knee and ankle joints, during the complete gait cycle. This is done by means of two elastic actuators whose elastic constants are changed for the different phases of the cycle to approach a normal profile. This concept aims at improving the gait pattern, avoiding the need to swing the trunk and the requirement of a high energy source.

This paper presents the design of the actuators in this system. The principle under the design is to use biomechanical data of the leg to determine the configuration of the actuators and actions applied on both joints. In Section 2, main problems of pathological conditions in patients with muscle weakness are explained in order to define the requirements of our actuator system. Next, an analysis of the normal gait cycle is done to establish the conditions of actions to apply on both joints. Based on this study an actuator system for knee and ankle was developed and its design will be detailed in Section 4. Finally, results and conclusions are given.

Fig. 1. Example of a typical commercial knee–ankle–foot orthosis, formed by three lateral bars (thigh, shank and foot) connected by two hinges.

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2. Pathological gait The absence of the necessary muscle control on the leg segments results in different impacts in locomotion, which can go from a non-desired gait pattern to the most severe, a collapsing body. For the design of an actuator system it is interesting to analyze what the situation can be in each joint when these problems occur [15].

2.1. Knee joint Knee extensors are active in both stance and swing phase. Their role is to control the rate of knee flexion, induced by the ground reaction force in stance and by the inertia of leg movement in swing.

In quadriceps weakness the situation is that, in the end of stance phase, the knee joint flexes in an uncontrolled manner, putting the patient at risk of falling. Besides, extension of the leg at the end of swing phase is also compromised, resulting in non-desired movements in a normal gait pattern. The other leg muscles (hamstrings or calf muscles) are less capable to accomplish the needed functional knee control.

A strategy to overcome the absence of knee extensor power and avoid uncontrolled knee flexion during stance is to prevent the ground reaction force from passing behind the knee joint. This may be achieved by increased hip extensor activity and anterior trunk bending (which displaces the centre of gravity of the body forwards). A possible long-term consequence of this condition is hyperextension of the knee due to the continued thrusting of the joint into full extension for security. If hyperextension develops, the knee will become more inherently stable due to its posterior displacement relative to the line of the ground reaction force during stance phase. However, the increased extension moment will induce a further hyperextension deformity such that this secondary problem may become a major clinical concern.

2.2. Ankle joint Ankle plantar flexors are active during stance and swing phases as well, and their functions are to control dorsiflexion in early stance phase and to provide active push off in late stance phase. In the case of ankle dorsiflexors the value of dorsiflexion moment in the ankle during gait cycle is much lower than that of plantar flexion, and its functions are to avoid uncontrolled plantar flexion after heel strike and to maintain the foot in the right position to avoid contact with the ground during swing phase.

Muscle weakness around the ankle joint may also result from pathological conditions such as poliomyelitis. These pathologies affecting the foot and ankle alone will seldom immobilize a patient totally; however, the resultant loss of joint control and the consequent compensatory maneuvers will usually result in a hazardous, less efficient is more energy-consuming gait. The most common problems present in these patients are the absence of control of plantar flexion after heel strike (foot flat) and the contact of the foot with the ground during swing phase (drop foot). Besides, absence of active push off at late stance phase is observed, but they usually adopt strategies to overcome this problem.

3. Functional analysis of gait as inspiration The requirements for an actuator system can be established in terms of the intended function the knee and ankle orthotic joints should feature in order to approach a normal gait. If an external actuator system is going to be designed for this purpose, an analysis of the functional actions of the musculo-skeletal system from mechanical point of view during the gait cycle might be a guide to understand the way to act on the lower limb joints with a portable system, trying to imitate this functionality [19,20].

3.1. Knee joint Analyzing the average gait data at natural cadence [16], the relationships between gait variables and energetic transformations during the cycle are studied in order to understand the mechanical basis involved in lower limb joints performance.

All data are normalized to subject’s weight.

It can be seen that approximately between 5% and 15% of the gait cycle (at stance), when the joint is absorbing the impact (through the action of quadriceps), the performance of the joint, by comparison of the angular position against the (normalized) torque at the knee–load–displacement relationship –, is similar to an elastic performance, as this relation is approximately linear.

During this interval the musculo-skeletal activity is dedicated to power absorption and then immediately, corresponding to the extension recovery, this power is recovered until approximately 30% of the phase. This roughly elastic relation starts almost at the beginning of the gait cycle at 1.4% and is maintained after the energy generation zone, until approximately 50% of the cycle, when the knee extension is completed and the knee is ready to swing.

The evolution of the ratio between torque and angle for the knee joint is illustrated in Fig. 2 (normalized data). Various different modes of operation can be identified by analyzing how the rotational displacement of the joint relates to its torque in the gait cycle.

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Fig. 2. Normal gait average biomechanical data. Knee angle vs. torque in the sagital plane (complete gait cycle) and optimal elastic adjustment.

According to the intervals defined in Fig. 2, the shock absorption area and the recovery of extension just before the swing phase is characterized by the segment A–B–C. The flexion phase during swing is characterized by the segment C–D and the extension phase during swing is characterized by the segment D–E. 3.2. Ankle joint A similar characterization process can be done to analyze the ankle joint. Winter’s data from normal subjects is used to analyze the functional characteristics of the joint during the gait cycle. The evolution of the ratio between torque and angle for the ankle joint is illustrated in Fig. 3 (all graphs present normalized data). Three intervals can be discriminated with a characteristic biomechanical behaviour. Between 0% and 5% (interval 1) when the joint is controlling the plantar flexion its behaviour can also be assumed as elastic and approached to an elastic linear model. Between 5% and 45% approximately, while dorsiflexion takes places during stance, a power storage phase is present also with a linear elastic approach. At interval 3, during plantar flexion at late stance, coinciding with the highest power energy demand, another nearby elastic relation can be found.

After interval 3 there is almost no action in the joint, coinciding with the displacement of the foot during swing, only little torque and power is required to produce the necessary dorsiflexion to avoid contact with the ground.

4. Actuator design The actuator system presented in this paper is inspired by the function of the real muscles of the leg, and it is conceived to mimic its behaviour. Consequently the starting point in the design process will be the actions in the joints presented before.

From an engineering perspective an analogy of the operation of the human musculo-skeletal system of the lower limb with a mechanical system can be established and its functionality can be considered as actuation functionality or as a combination of them. In this section, actuation functions for each joint are identified during each phase, and thus, a first approach of a mechanical adjustment of gait by elastic means is described. Finally, a design of an actuator system is proposed.

This prototype was adapted to the GAIT orthosis (Fig. 4), a novel knee, ankle and foot orthosis (KAFO) [17]. It has been designed to be modular and adaptable to subjects with different anthropometric characteristics. The mechanical structure is formed by a single sided frame with two joints, knee and ankle. Knee hinge is performed by a four-bar mechanism to follow the displacement of the helicoidal instant axis of the knee in the lateral side. Ankle hinge is performed by a single hinge placed on the malleolli.

The main difference with other systems is the knee joint concept used [18]. A four-bar mechanism was designed to follow the movement of the human knee, simulating the movement of the knee cruciate ligaments. The performance of this joint can be considered as a rotation around an instant helicoidal axis, which changes its position during the displacement. The

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Fig. 3. Normal gait average biomechanical data. Ankle angle vs. torque in the sagital plane (complete gait cycle) and optimal elastic adjustment.

Fig. 4. GAIT orthosis.

four-bar hinge therefore follows the displacement of said helicoidal instant axis of the knee in the lateral side, to get a movement similar to the physiological movement of the knee. A genetic algorithm methodology was used to adapt the lengths and angles of the bars in the hinge, as presented by Baydal et al. in [18].

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Fig. 5. Knee actuator concept and functionalities during the gait cycle. Stacked discs constitute the stance control spring (top) and one compression spring (below) for the swing control.

This new concept provides a reduction on the displacements of the orthosis due to misalignments of the orthotic joint with respect to the human joint. The physiological ankle joint acts as a single hinge and it was implemented as a monocentric hinge with its centre of rotation in the middle of the distal tips of the two malleolli.

Based on the functions of the muscles in the leg during gait cycle, functions of an actuator system can be defined as follow: Knee Stance phase: Shock absorption and knee assistance back to full extension.

Swing phase: Free knee flexion in early swing and assisted knee extension in late swing to prepare for the next foot contact.

Ankle Stance phase: Control of plantar flexion after heel strike to prevent foot flat.

Swing phase: Active push off in early swing and control of plantar flexion to prevent drop foot.

For the actuator design we will consider the worst possible scenario, where there are no residual functionalities on knee and ankle, and the actuators should provide the same actions as the real muscles. Due to the different characteristics of each joint two, independent actuators have been designed to fulfil the required actions.

4.1. Knee joint The concept for the knee actuator during the gait cycle is described in Fig. 5.1 During the start of the shock absorption (A) and weight bearing (B), the stance control spring (green) is active and compresses while the leg is in the ground. At start of the swing phase (C), the swing control spring (blue) is enabled and the stance control spring is disabled. The swing phase spring compresses and recovers the energy completing the swing phase.

An electromagnetic solenoid is used to switch between springs as a function of the gait cycle and to lock the knee flexion if required. Both springs will provide the required action for each part of the cycle when compressed (by the user’s weight during weight bearing, and by a combination of inertia and push off action of the ankle during swing phase). The change between springs will be done with the leg completely extended.

1

For the interpretation of color in this figure, the reader is referred to the Web version of this article.

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For the adjustment of the linear elastic constants of the actuator springs we need to calculate first the theoretical torsional elastic constant of an action based on the joint angle. Let us consider the intervals defined in previous section: Interval 1. Shock absorption–recovery of extension: We can consider the interval between the beginning and 50% of the GAIT cycle (stance phase) as an approximately linear relation between applied torque and flexion angle, as can be seen in the torque–angle diagram shown in Fig. 2. This zone corresponds to the interval between points A and C. The interval 1 features a practically constant ratio, similar to the elastic constant defined by Hook’s law, which holds proportionality between mechanical stress and strain, so a line equation can be adjusted to this interval with a small error.

Due to the actuator concept, where there is no action on the knee when it is completely extended, the optimal adjustment is not practically feasible. It would be required to have two different neutral positions (with no action on the joint), for stance and swing, with different values of knee flexion. Instead of it, the selected solution was to adjust a line from point 0-0 (Fig. 2) to the maximum torque value, point 22–0.62 in the angle–torque diagram. From the equation of this line a torsional elastic constant can be obtained: T ¼ K 1 astance ;

ð1Þ

where T is the joint torque in N m/kg, astance the joint angle for stance phase in degrees, and K1 the torsional elastic constant in N m/kg deg.

Intervals 2 and 3. Flexion and extension during the swing phase: Interval 2 presents a non-linearity in the torque–displacement curve that can be referred to as pseudoelasticity. This behaviour is present when, after reaching given loading stress, the deformation strain augments considerably with minimum applied stress. A superelasticity phenomenon is known for its non-linearity during unloading. In the case of the knee joint, the torque–angle relationship holds a pseudoelasticity behaviour during loading. Another interval featuring an approximately linear relation between torque and angular displacement can be considered between points D and E, interval 3.

The optimal adjustment for both intervals can be a line between A and D, but the need of starting in the stance position, imposes point 0–0, as the starting point. The obtained adjustment is depicted. The torsional elastic constant is obtained from the expression: T ¼ K 2 aswing ;

ð2Þ

where T is the joint torque (referred to user weight) in N m/kg, aswing the joint angle for swing phase in degrees, and K2 the torsional elastic constant in N m/kg deg.

Multiplying these constants by the angle in the knee during the complete gait cycle we can obtain the action of two theoretical knee actuators featuring this elastic performance. Fig. 6 shows the elastic adjustment for both constants and the theoretical action of the actuator with the configuration explained before, in comparison with average data of healthy subjects.

This action is a combination of both adjustments. It uses K1 for stance angles and K2 for swing angles.

With the definitions of the separate actions per each gait cycle, the next step was to develop an actuator prototype to test the concept. It was decided to use a linear actuator placed in the sagital plane. This actuator will apply force on the orthosis, generating the needed moment during gait cycle.

Additional requirements considered during the design process were:

Stance range of movement: 0–22°; Wing range of movement: 0–65°; Maximum range of flexion: 95°; Minimum torque values for 90° of knee flexion (sitting position).

The solution adopted for the last requirement was to align the force applied by the actuator with the centre of rotation of the orthotic joint, making minimum the applied moment. From the rest of the requirements and after considering different geometric configurations, the solution adopted for the knee actuator is formed by two telescopic cylinders, one containing the stance spring and the other containing the swing spring, following the concept explained in Fig. 5.

Two types of springs were used in the prototype. Swing springs are compression springs, made of stainless steel and with right hand direction of the helix. Different constants were available for selection. An elastic element (8 mm length) was used for the stance control, applying K1 at the joint with a maximum longitudinal stroke, corresponding with a flexion limit. A compression elastic element (57 mm length) is provided for the swing phase. Compression springs made of Stainless Steel Type 302, were used.

Selection of stance and swing springs in the knee actuator is done as follows: assuming no residual actions on the joints, from equations of elastic adjustment done previously and with patient weight data, maximum torque values (stance and swing) were obtained: T ¼ K 1;2 astance=swing Patient weight: Torque provided by the actuator comes from the expression

ð3Þ

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Fig. 6. Elastic adjustment for both elastic constants during the complete gait cycle and final actuator action on the knee.

T ¼ F actuator ðx cos a þ y sin aÞ

ð4Þ

applied at maximum stance and swing flexion angles. While the orthotic knee joint is a four-bar mechanism it is necessary to calculate the instant centre of rotation at those maximums in order to obtain x and y distances. This has been done finding the intersecting point of both central bars of the mechanism. The expression of a linear spring gives us the linear elastic constant of the springs: F actuator ¼ K spring Dlactuator :

ð5Þ

The action provided by the actuator with the obtained constants for both springs is different to that calculated in the elastic adjustment. This is due to the relative displacement of the actuator in the orthosis, but the differences are not significant.

4.2. Ankle joint The concept of the ankle actuator differs slightly from that for the knee. It is formed by two springs, but, due to the characteristics of the joint actions, it is not mandatory to change between springs when the joint goes into flexion during different phases of the cycle. Considering dorsiflexion and plantar flexion separately, one of the springs is active only in positive

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angles (dorsiflexion) and the other one in negative angles (plantar flexion). With this concept one of these springs will control plantar flexion after heel contact, and the other one will be active from middle to late stance phase, controlling dorsiflexion. During swing phase the foot will stay in neutral position, restricted by the stiffness of the spring controlling plantar flexion, avoiding contact with the ground.

Considering the defined intervals it is possible to adjust the elastic constants: Interval 1. Plantar flexion control at the start of stance phase: Let us consider interval 1 defined in Fig. 3. We can consider the relation between torque and angle as almost linear in the interval. Notice that torque during this interval is very low.

Intervals 2 and 3. Dorsiflexion in the first half of stance phase and plantar flexion at late stance phase: In this case both intervals were approximated by a linear relation. The adjustment done in this case approximates the torque–angle behaviour by line O–D (Fig. 3). The adjustment to the torque–angle curve is not optimal, but it is imposed by the concept of the actuator, which has two separated actions depending on the sign of the joint displacement.

Angular elastic constants for both intervals are obtained from the expression of the adjustment: T ¼ K 3;4 aankle ;

ð6Þ

where T is the joint torque (referred to user weight) in N m/kg, aankle the joint angle in degrees, and K3,4 the torsional elastic constants in N m/kg deg.

Theoretical action of an actuator with these constants is obtained by multiplying both constants by the ankle angle during gait cycle. Adjustments can also be seen in Fig. 3.

Selecting the right constant for each interval (K3 for dorsiflexion and K4 for plantar flexion) we obtain the actions required from the actuator. Fig. 7 illustrates the theoretical result of these actions.

The prototype designed is a linear actuator which applies force in the sagital plane, resulting in the required moment on the joint.

Requirements for the ankle actuator were Range of movement going from 10° to 20°.

Actions (moments) in other axis different to that of ankle flexion have to be minimum.

After evaluating different geometric configurations it was decided to use the concept illustrated in Fig. 8. It is formed by two cylinders, each containing one spring, and an internal part which can slide in two directions. When this part enters the cylinder one of the springs is compressed, and when the part slides out of the cylinder it is the secondary which is compressed. Actions of each spring are independent.

The selected springs are formed by small discs. The actuator attached to the orthosis is shown in Fig. 8.

Procedure for the calculation of both linear elastic constants is as follows. From equations of elastic adjustment and with patient weight data, maximum torque value is obtained: T ¼ K 3;4 aankle Patient weight:

Fig. 7. Elastic adjustment for both elastic constants during the complete gait cycle and final actuator action on the ankle.

ð7Þ

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Fig. 8. Detail of the ankle actuator and actions applied on the ankle: (a) dorsiflexion spring (K3, proximal cylinder) is active at initial stance, (b) the controlling spring (K4) is active during plantar flexion, preventing the foot fall.

Torque provided by the actuator comes from the expression T ¼ F actuator ðx cos a y sin aÞ:

ð8Þ

The expression of a linear spring gives us the linear elastic constants: F actuator ¼ K spring Dlactuator :

ð5Þ

The actuator assembly is 30 mm in width. The stance control spring is formed by small discs, which can be grouped in different ways to get different constants. Belleville Washers from 300 Series Stainless Steel were used for this actuator. Each stacked spring is 8 mm in length and the assembly is 30 mm wide.

It is important to consider that the calculated actions for both actuators are based on the supposition that the joint angles will be the same during gait cycle for a patient wearing the system than those of a normal subject. This has proved to be a good approximation as a starting point for the design of the actuators.

5. Experimental evaluation The designed actuating system was tested with two poliomyelitis patients. To drive the switching mechanism to change between springs in the knee actuator, two systems were used. The first one was a cable controlled by the ankle angle. This cable pulled the switching mechanism of the knee actuator when ankle angle reached its maximum of dorsiflexion, releasing the knee. The second one was a solenoid controlled using signals from shank velocity, knee angle, and actuator status.

To evaluate the performance of the system a set of sensors was attached to the orthosis to measure different variables.

These sensors were: A goniometer on the knee: A potentiometer attached to the orthotic joint was used to measure the angular displacement.

Two gyroscopes: One measuring angular velocity of the shank and a second measuring angular velocity of the foot.

Initially, constants of the springs were calculated using the previously described method to provide actions comparable to those of a real subject.

The two poliomyelitis patients were both left leg affected, but with different severity.

The first patient was a 44-years-old woman with poliomyelitis on her left leg (Fig. 9). She uses her own orthosis, which is a double bar knee locked system. It has limited movement in the ankle joint. She had experience with other orthotic systems with knee lock/unlock system, as the UTX [4], for that reason she rapidly learnt how to use the Gait system. Other characteristics of the patient were weight, 78 kg, and height, 150 cm.

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Fig. 9. Picture of the actuated orthosis worn by a poliomyelitis patient before (left) and after (right) the orthosis has been fitted to the patient with additional padding (foam plastic).

Initially, the patient was asked to start walking in a walking bridge with bars on both sides to get confident with the new orthosis. To tune the knee lock/unlock mechanism properly the patient walked through the bars back and forth for a number of times. It took approximately half an hour before the patient started walking with only two crutches.

The second patient was a 46-years-old man affected with poliomyelitis in his left leg. He had also affected the right leg although with less gravity. The subject weighed 76 kg and was 172 cm height. His gait was autonomous with the help of crutches to walk and he did not use any kind of orthotic device. Mobility in the ankle joint of the patient was very low, having a maximum of dorsiflexion of very few degrees. This caused a bad performance of the cable controlled system. For this reason data studied will be only that collected with the solenoid controlled system. He started walking with a crutch, but he fast gained confidence and, after half an hour he was able to walk with no help. In both cases, before starting the tests, it was necessary to add some material (foam plastic) between patient’s leg and the orthosis, due to shape and size of the leg.

6. Results and discussion Example of data collected during one trial is illustrated in Fig. 10. The gait pattern for each cycle can be identified. This pattern can be observed in the three signals, and it is quite constant during the trial, despite some differences in the maximum knee angle.

Fig. 10. Resulting kinematics of patient testing: example of a walking trial (detailed).

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A first conclusion is that the switching mechanism works properly, allowing flexion of the knee during swing and providing stability during stance. On the other hand it can be observed that knee flexion during stance is almost non-existent. Possible causes for this could be excessive stiffness of the stance spring in the knee actuator or lack of patient confidence in the system.

Stance phase: In the case of patient 1 during stance phase there is no knee flexion. Observation of videos recorded during testing showed that the patient was not completely relying on the orthosis during stance. Instead, she was partially leaning on the crutches and did not compress the stance spring.

The foot velocity data features a slight increase during all first half of the phase (area 1). The fact that these values are not very high and ankle movement is not sudden indicates that plantar flexion after heel strike (foot flat) is being controlled by the ankle actuator.

Swing phase: During first half of swing phase (0–40 deg) knee flexion increases in two intervals. The primal is produced immediately after releasing the knee. It can be observed that a certain degree of positive angular displacement of both, shank and foot. This can be an indicator of push-off action from the ankle actuator, which causes knee flexion and an elevation of the heel. Then, knee flexion goes to 40 deg due to inertia caused by hip extension. Another increase in shank and foot positive velocity can be observed.

During second half of swing phase (40–0 deg) the knee extends completely and velocity of shank and foot goes into negative values. This extension action on the knee is caused by a combination of inertia and knee actuator action. Different tests done at different speeds showed an important influence of gait speed in this part of the cycle. At slow speed the extension action on the knee during swing with the selected spring was insufficient and extension of the leg was not complete.

Although the self-selected speed was low, the assistance to extension was clear while being able to complete full extension at the end of the swing phase, as has been found from the mean values of peak knee flexion. After a brief period of time, the subject learnt to ensure the hinge lock during initial stance, relying on the device. From examination of the ankle angle, the excessive dorsiflexion with the subjects own orthosis, progresses in an adequate manner during the stance phase as compensated with the ankle actuator. The plantar fall is limited at maximum 5 deg. It was also found that the orthosis successfully avoids the need of the compensation by trunk movements. In the first group of walking trials the subject did not succeeded in releasing the knee hinge in 20% of the situations. In the second group of tests, the confidence was incremented with no unsuccessful situations (locked knee during stance). The self-learnt strategy by the patient consisted in maintaining full extension during stance, aided with a crane, which not allowed us to quantify the loading response by the orthosis.

Another conclusion can be extracted from the fact that, during swing phase, when the foot is not in contact with the ground (areas 3 and 4), foot and shank velocities are not equal, this being an indication of relative displacement between them. If this displacement is high the foot can touch the ground (drop foot). This difference in velocities can be explained by a low stiffness of the second spring in this case, which compresses during early swing due to inertia, and returns its energy in late swing, making the foot turn in relation with the shank.

7. Conclusions This paper presents the design of an actuator system for knee–ankle–foot orthosis based on biomechanical data. This actuator system was integrated in a KAFO prototype for its validation with patients. The experimental results demonstrate the feasibility of significantly improving gait in patients with proximal leg weakness by means of compensations applied by an actuated orthosis.

The assistance to the extension of the leg in late swing phase is remarkable, as an important improvement in comparison to other systems and according to the opinion of the users. Our design hypothesis for selection of the springs has been shown in preliminary cases as a good approximation for the knee actuator: ranges of joint angles during compensated gait in patients are comparable to the normative average data. Nevertheless, the calculated values for the springs obtained in the case of the ankle joint are higher and more research has to be done in this field. It is clear that the concept of a linear elastic actuator based on the joint angle which stores energy to release it when required is valid, but the required actions to apply on the joints depend on many factors, e.g., walking speed, pathology, etc.

The knee actuator concept should be revised. It would be interesting to apply not only extension actions on the knee, but also to incorporate flexion action. A concept based on torsion springs in combination with electric actuators could be compatible in order to provide such actions. In the case of the ankle, performance of the actuator avoiding drop foot and foot flat was appropriate. It is not so clear that it was applying active push-off, prior to the swing, observed actions can be due to remaining muscles or patient’s walking strategies. The major problems observed here were the difficulty to apply actions with such small displacement (movement of the ankle restricted), and the influence of the stiffness of other elements (in this case, an in-shoe insole) in the applied actions. Each pathological case presents particular characteristics and thus, the mechanical fitting and control system tuning procedures require customization of the solution. Future work will consist on the clinical evaluation of mid-term and long-term effects of our functional compensation concept, and also continuing the detailed analysis of the performance of the actuators during different conditions.

In general, previous studies have been focused on providing a free swing-phase motion while preventing the knee flexion during stance [5]. The obtained knee joint angle profiles seen in this study might imply the ability to acquire a much more physiological gait preventing from the lateral movement, very common in these kinds of patients. Currently, we are exam-

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ining the situation during the swing phase, and in particular, the effects of the spring for extension assistance during the initiation of the oscillation and prior to heel contact.

Acknowledgements The work presented in this paper has been partially funded through Grant IST-2001-37751 of the European Commission.

The authors thank Ossur hf and Roessingh Research and Development for their contributions in materials and experimentation.

The work presented in this paper has been partially funded by the European Commission.

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