Laser Surface Treatment of Metal Implants: A Review Article

www.me‐journal.org Journal of Metallurgical Engineering (ME) Volume 4, 2015 doi: 10.14355/me.2015.04.006

Laser Surface Treatment of Metal Implants: A Review Article Ibrahim M. Ghayad, Nabil N. Girgis, Wafaa A. Ghanem Central Metallurgical Research & Development Institute (CMRDI), P.O.Box:87 Helwan, Cairo, Egypt ighayad@yahoo.com

Abstract Metal implants are mainly fabricated from stainless steels, cobalt base or titanium base alloys. In addition to mechanical properties, metal implants must have excellent corrosion resistance as well as good bioconductivity and biocompatibity.

Biocompatibity is achieved on a metal implant by the application of proper surface treatment. This paper reviews the techniques used for surface treatment of metal implants with a special focusing on laser surface treatment as a new and promising technology for surface treatments. Published papers indicated that surface treatments performed on metal implants are mainly of two types; the first concerned with the deposition of hydroxyapatite (HA), bone‐like material while the second focused on the deposition of diamond‐like carbon (DLC). Present review covers in detail the laser systems used for either HA or DLC coating.

Keywords Laser; Metal Implants; Diamond Like Carbon; Hydroxyapatite

Introduction Metallic materials are widely used as internal fixation implants in bone fracture surgery to provide temporary support during the bone healing period [1] Stainless steels, titanium alloys, cobalt‐chromium based alloys and magnesium alloys are commonly used for this purpose. A limitation of these alloys is the possible release of toxic metallic ions and/or particles through corrosion or wear processes [2‐6] that lead to inflammatory cascades which reduce biocompatibility, causing tissue loss or adversely affect their mechanical integrity before healing [7‐13].

Several techniques have existed to improve the corrosion resistance and biocompatibility of above mentioned alloys; the most effective is surface treatment. In this text two types of surface treatment, used for metal implants, will be discussed. The first concerns with the deposition of phosphates, mainly hydroxy apatite (HA), while the second deals with the deposition of diamond‐like carbon (DLC).

Hydroxylapatite (HA), is currently used as a biomedical material due to its excellent biocompatibility and bioactivity which attributed to its chemical and structural similarities to bone and tooth mineral [14]. On account of its low strength and high brittleness, HA is used as a bioactive coating on metallic substrates.

When natural human heart valves undergo degenerative process such as calcification of leaflets, due to infections, ageing, or dietary problems, artificial heart valves replacement is a solution and low temperature isotropic pyrolytic carbon (LTIC) is the most common and widely used material. Unfortunately, its blood compatibility is still not adequate and as a result, patients must continue to take blood anticoagulation medicine. However, amorphous carbon (a‐c) or diamond‐like carbon (DLC) films are potential biomedical materials due to their chemical inertness, impermeability and excellent mechanical properties.

Many methods have been developed to prepare HA coatings and to deposit DLC thin films. These methods are: sol‐gel process , electrophortic deposition , the sputtering process, biomimetic methods plazma spraying , physical vapor deposition as well as laser techniques [15‐21].

This review focuses on laser surface treatment as a new and promising technology for surface treatments of metal implants. Hydroxyapatite coatings and diamond‐like carbon films deposited on metallic substrates using laser technology are described and discussed.

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Types of Laser Surface Treatment Techniques Krishnan et al. [21] reviewed novel laser techniques used for synthesis of nanostructured coatings and thin films.

These techniques are: Laser pyrolysis, laser assisted atmospheric chemical vapor deposition (LAACVD) and pulsed laser deposition (PLD). In the laser pyrolysis technique the aerosols generated using an ultrasonic particle generator or compressed air nebulizer, (metal nitrate solution droplets), are passed through a quartz tube. The laser was coupled into the quartz tube in a coaxial manner and the laser beam along with the aerosol was directed towards the substrate kept very close to the focal point of the focusing lens. The laser interacts with the aerosol during its transit through the tube as well as at the substrate, leading to reaction with atmospheric oxygen at the substrate forming an oxide coating. By adjusting the process parameters, localized novel nanostructures are obtained. These localized structures are produced because of the Gaussian distribution of laser power coupled with non‐equilibrium nature of surface reaction.

In laser assisted atmospheric chemical vapor deposition (CVD) technique the coating precursor used is a suitable metal organic compound dissolved in ethanol. The precursor is atomized using an ultrasonic particle generator to produce aerosols of 1 to 5 micron size. Upon laser‐precursor aerosol interaction, the metal organic compound undergoes vaporization without decomposition and the vapors react with the oxygen present in the atmosphere at the laser focused spot on the substrate surface to give oxide coating. Uniform large surface area coatings are obtained by rastering. A computerized raster stage is used for this purpose.

Pulsed laser deposition (PLD) is a thin film deposition technique where a high power pulsed laser beam is focused inside a vacuum chamber to strike a target of the desired composition. Material is then vaporized from the target and deposited as a thin film on a substrate. This process can occur in ultra high vacuum or in the presence of a background gas, such as nitrogen which is used when depositing nitrides. When the laser pulse is absorbed by the target, energy is first converted to electronic excitation and then into thermal, chemical and mechanical energy resulting in evaporation, ablation, plasma formation and even exfoliation. The ejected species expand into the surrounding vacuum in the form of a plume containing many energetic species including atoms, molecules, electrons, ions, clusters, particulates and molten globules, before depositing on the typically hot substrate.

Laser surface melting technique modifies the metal implant surface by refining the structure after melting and rapid crystallization. It is well‐known that grain refinement improves mechanical properties and corrosion resistance.

Hydroxiapatite Coatings Nelea et al. [22] investigated hydroxiapatite thin films grown by pulsed laser deposition (PLD) (Figure 1). The targets were prepared from high‐purity (99.98%) polycrystalline HA powder (grains 100 mm) by cold pressing followed by annealing in air at 600oC. A Ca/P atomic ratio of 1.6 was inferred by energy dispersive X‐ray spectrometry. The TiN buffer interlayer was deposited by reactive pulsed laser deposition (RPLD) by ablating a stoichiometric TiN target in nitrogen, whereas the metallic substrate was heated at 650oC. A Series of 104 subsequent laser pulses were applied for the growth ofʹ < 1 mm thick HA film. The average ablation rate was 0.5 A.U. per pulse.

In the work of Grigorescui et al [23], a hydroxyapatite target was ablated in an oxygen controlled atmosphere chamber by means of a KrF excimer laser beam of 248 nm wavelength, generating pulses of 10 ns duration. The laser source was operated at a frequency repetition rate of 2 Hz. The spot area was between 4.5‐5 mm2, and the incident laser fluence on the target was 2.6 J/cm2. In order to avoid target piercing during multipulse ablation, a rotation movement was applied. The coatings were deposited on 15 mm diameter disks made of high purity c.p. Ti, heated at 400°C. Prior to every film deposition, the chamber was evacuated down to a residual pressure of 10‐4 Pa. Then a dynamical flux of 10 Pa O2 was introduced and carefully monitored during deposition. For obtaining one film, 15000 subsequent laser pulses were applied. After deposition, the films were treated in water vapour enriched atmosphere at 400°C for 6 hours, in order to improve their crystalline state. Other samples have been prepared by the insertion of a buffer layer interlayers of TiN deposited by PLD before the application of HA coatings. The thickness of TiN layer was estimated to ~ 200 nm, while the thickness of HA layers was estimated to be in the range 1‐1.5 μm.

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In the work of Syed et al. [24], a powder, blend of different powders or powders from two or more powder feeders were delivered to a laser generated melt pool on a substrate surface (Figure 1). The relative movement of the laser and the substrate causes the enlarged melt pool to solidify and a well‐bonded clad/tack with the substrate is formed. This method can also produce functionally graded materials (FGMs). Depending upon the requirement for the final part, varying proportions of different materials can be added on top of each other to obtain a graded structure. The main purpose of using a graded‐layer coating is to overcome the stress‐related failure of composite coatings that might result from a sharp contrast of properties, such as thermal expansion behavior.

FIGURE 1 SCHEMATIC DIAGRAM OF THE EXPERIMENTAL APPARATUS FOR POWDER‐POWDER DEPOSITION (SCHEMATIC)

Vasanthan, Drukteinis and Lacefield (2008) [25] deposited commercially pure hydroxiapatite on cp‐Ti substrate using pulsed laser technique. They used KrF excimer laser (248 nm), with an energy density of 4‐10 J/cm2 and a repetition rate of 30 Hz while the substrate temperature was kept between 625 oC and 715oC. A deflecting mirror was used to guide the laser into a controlled atmosphere chamber where it ablated the HA target to generate a plume of atoms, ions, and particles (Figure 2). These particles then coalesced and deposited themselves onto the surface of the heated, rotating alloy in an argon/water atmosphere, thereby forming the coating.

FIGURE 2 SCHEMATIC DIAGRAM OF THE PULSED LASER COATING SYSTEM

Rau et al [26] deposited fluorinated hydroxyapatite (FHA) films on Ti using pulsed laser technique. The use of FHA in biomedical applications is expected to be advantageous, since it can promote mineralization and crystallization of calcium phosphate during the dental and bone formation processes and ensure the formation of a mechanically and functionally stronger bone by providing the fluorine release at a controlled rate. The synthesis of fluorinated hydroxyapatite was carried out by mixing CaO, (NH4)2HPO4 and NH4F. FHA films were deposited on heated Ti substrates in a high vacuum PLD chamber. Depositions were performed by ablating the sintered FHA rotating target with a laser beam generated by a pulsed KrF excimer‐laser (l = 248 nm); the laser pulse duration was 17 nm, and the repetition rate was 5 Hz. The laser beam was oriented with an inclination angle of 45o with respect to the target, whereas the substrate and target were assembled in a frontal geometry at 4 cm of reciprocal distance.

The PLD chamber was evacuated down to a base pressure of 1x10‐6 mbar prior to the film deposition; then,

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depositions were performed at 5x 10‐4 mbar in a controlled dynamic pressure produced by the N2 gas flow, introduced directly into the chamber through a needle valve. The deposition time for each sample was 25 min, for a total of 7500 pulses. A set of FHA films were deposited at fixed temperature (400oC) and different energy of the laser beam (2, 3, 5, 7 J/cm2).

Bai et al [27] reported processing and characterization of functionally graded hydroxyapatite (FGHA) coatings incorporated with Ag as an antibacterial component. FGHA‐coated implants incorporated with antimicrobial agents are able to prevent or cure infections by releasing directly the antimicrobial agents to local regions. The amorphous top layer of the coating allows a higher release rate of silver due to its higher dissolution immediately after implantation; the crystalline layer will maintain the silver as a reservoir, resulting in long term infection protection.

Man, Chiu and Guo [28] studied the effects of laser drilled micro‐holes at the surface of the metals in order to achieve a strong adhesion for a structural joint or a bone tissue fixation for medical implants. The effect of the number of holes per unit area on the adhesion strength of the adhesion joint was evaluated. Results showed that the number of holes per unit area on the adherent surface logarithmically correlated with the bonding strength.

Other holes geometries are suggested for enhanced adhesion and bone tissue fixation.

In the work of Hu et al. [29], the In‐situ synthesis and fabrication of tricalcium phosphate bioceramic coating on commercially pure Ti by laser rapid forming (LRF) was investigated (Figure 3). The powders of CaCO3 and CaHPO4∙2H2O were used as raw materials. The powders were mixed well to yield a powder with a Ca/P ratio of 1.5 and then loaded into the powder feeder. The cp‐Ti substrate plate was cleaned with absolute alcohol and acetone to remove surface contamination and then fixed to the NC working table. A computer controlled continuous wave CO2 laser was used to scan the surface of the substrate with the energy density of 60 J/mm2 (laser power of 400 W; scanning velocity of 100 mm/min; laser spot diameter of 4.0 mm) and a powder feed rate of 3.0 g/min. The coating was fabricated in a closed chamber filled with argon (Ar) gas to avoid oxidation, and Ar gas was also chosen as the powder carrier gas. In the LRF system, the NC working head controlled the movement of the laser focusing mirror and the lateral powder feed nozzle so that the laser spot and powder delivery site moved synchronously. Then the mixed powder was melted and solidified rapidly with the laser spot moving away, so a single track of coating was fabricated. Since the working head moved in cycles with an overlap ratio of 35% between adjacent tracks, multiple tracks of coating eventually converged to form a complete coating. The phase composition of the coating contained 95 wt.% of β‐TCP and 5 wt.% of α‐TCP. Three layers were found in the coating: a ceramic layer, a transitional layer, and the substrate layer. In the transitional layer, interpenetration of phases was observed. The bonding strength between the coating and the cp‐Ti substrate was in excess of 40.17 MPa. Furthermore, the static immersion test has confirmed that the coating not only prevented the corrosion of cp‐Ti, but also induced the redeposition of β‐TCP in synthetic saliva.

FIGURE 3 SCHEMATIC PRESENTATION OF THE LASER RAPID FORMING PROCESS

Diamond-like Carbon Coatings Diamond like carbon (DLC) coatings were performed only under vacuum using pulsed laser technique. Authors used ArF, KrF or Nd‐YAG lasers.

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Wang et al. [30] succeeded to deposit diamond‐like carbon films containing metal elements on biomedical Ti alloys.

In their work, a series of amorphous hydrogenated carbon films (with and without metal nitride interlayer) were deposited on biomedical Ti alloy (Ti ‐6A1‐4V) substrates using a cathode arc evaporation (CAE) system.

The experimental apparatus of the pulsed laser ablation system for film deposition used by Xiong et al [31] is schematically shown in Figure 4. It consists of a stainless steel high vacuum chamber pumped by a turbo‐molecular pump with a base pressure of lxl0‐7 Torr. An ArF pulsed laser beam with the wavelength of 193 nm and the pulse duration of 21 ns (FWHM) is introduced into the chamber through an antireflection coated convex focus lens and a supersil quartz window. The stabilized output energy of 300 mJ/pulse and a repetition rate of 5 or 10 Hz are normally used. The laser beam impacts the target at an angle of 45o to the surface normal and has a spot size of 1‐2 mm2 at the target surface. The instantaneous power density at the target is estimated to be about 5x108 W/cm2. A 1 inch diameter pyrolytic graphite (PG) disk is used as an ablation target. It is kept rotating at 10 rpm during the process. The laser plume is emitted along the surface normal of the target, intensified within a few millimeters near the surface in a bluish‐white color, then expanding forward in a reddish cone. The unheated substrate is placed about one inch away and parallel to the target. A typical deposition rate of 0.1 nm/pulse is obtained, slightly varying with the deposition conditions, such as the vacuum condition and laser pulse energy. N‐type Si(100) wafers (resistivity 0.1 n‐cm), fused quartz slides, and cover glass slides are routinely used as substrates. The resulting films possess remarkable physical, optical and mechanical properties which are close to those of diamond and distinct from the graphite target used. The films have a mechanical hardness up to 38 GPa, an optical energy band gap of 2.6 eV and excellent thermal stability. Analysis of electron energy loss spectroscopy reveals the domination of diamond‐type tetrahedral bonding structure in the films with the spl bond fraction over 95 %.

Compared to other reported results of pulsed laser deposited diamond‐like carbon films, experimental results confirm that the laser wavelength or photon energy plays a crucial role in controlling the properties of pulsed laser deposited diamond‐like carbon films.

FIGURE 4 SCHEMATIC DIAGRAM OF THE PULSED EXCIMER LASER ABLATION SYSTEM FOR FILM DEPOSITION

Hanabusa and Tsujihara [32] deposited Diamond‐like carbon (DLC) films by laser ablation with the target made of frozen acetylene, instead of conventional graphite (Figure 5). They used a stainless steel vacuum chamber, 20 cm in diameter and 25 cm high, contained a quartz substrate on a heated holder. The target was prepared on a copper plate and placed in contact with a liquid nitrogen reservoir. Acetylene was blown through a quarter of an inch stainless steel pipe toward the copper plate. A rotary pump evacuated the chamber during the target preparation.

After a frozen acetylene layer grew into roughly 1‐2‐mm thick layers, the gas supply stopped and a turbo‐ molecular pump was turned on to improve the pressure of the chamber to 10‐6 Torr. It increased to 10‐5 Torr during deposition. The light source used was a 193‐nm ArF laser with a pulse width of roughly 14 ns. The repetition frequency was set at 10 Hz. The spot size of laser beams on the target was adjusted by a quartz lens with a focal length of 35 cm. Laser beam was moved constantly across the target to avoid ablation of the copper surface exposed after losing the frozen gas. The quartz plate was placed in the vacuum chamber at a distance of 40 mm from the target after surface cleaning. The C‐H bond was more abundant in the films deposited by the KrF layer.

It was possible to deposit particles‐free films in contrast to laser ablation of graphite targets. Experimental results suggest the importance of energetic and charged species ejected from frozen acetylene.

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FIGURE 5 SCHEMATIC DIAGRAM OF THE ARRANGEMENT LASER ABLATION FOR DEPOSITION OF DLC FILMS

In the work of Lackner et al. [33], Graphite targets were ablated with pulsed Nd:YAG laser at 1064 nm wavelength in argon and C2H2 atmospheres to deposit amorphous hydrogen‐free (a‐C) and hydrogenated (a‐C:H) DLC onto various steel substrates (AISI 1045H, B7, H13, D2, M2). The targets were rotated during the laser irradiation in order to avoid the formation of deep craters. Deposition rates between 15 and 35 nm/min have been reached for both a‐C and a‐C:H coatings, deposited in Ar and C2H2 atmospheres, respectively.

Honglertkongsakul, May and Paosawatyanyong [34] investigated the pulsed laser ablation of a graphite target by ArF excimer laser deposition at a laser wavelength of 193 nm and fluencies of 10 and 20 J/cm2 to produce diamond‐ like carbon (DLC) films (Figure 5). DLC films were deposited on silicon and quartz substrates under 1×10−6 Torr pressure at different temperatures from room temperature to 250°C. The effect of temperature on the electrical and optical properties of the DLC films was studied. Raman spectroscopy (LRS) showed that the DLC band slightly increased to higher frequency with increasing film deposition temperature. Spectroscopic ellipsometry (SE) and ultraviolet–visible absorption spectroscopy showed that the optical band gap of the DLC films was 0.8‐2 eV and decreased with increasing substrate temperature. These results were consistent with the electrical resistivity results, which gave values for the films in the range 1.0×104‐2.8×105 Ω cm and which also decreased with deposition temperature. We conclude that at higher substrate deposition temperatures, DLC films show increasing graphitic characteristics yielding lower electrical resistivity and a smaller optical band gap.

Yap and Tou [35] deposited diamond‐like carbon films (DLC) on silicon and glass substrates by using a 10 Hz, 4.7 ns, Nd‐YAG laser at third‐harmonic and fundamental wavelengths of 355 nm and 1064 nm, respectively. A pyrolytic graphite target was fixed in position while the laser beam, which was focused to a 0.5 mm diameter, ablated over an area 6.5 mm x 6.5 mm at an incident angle of 45° by an X‐Y stepper‐motor beam scanner. Laser fluencies of 21 J/cm2 and 42 J/cm2 were used in graphite ablation. The p‐type silicon wafer (100) with a resistivity of 10 ohm cm was used as substrates. In a typical deposition, substrate with dimension of 0.8 cm x 0.8 cm was placed at 3 cm distance from the target. The base pressure of the deposition was 10‐6 Torr.

In the study of Cho et al [36], functionally gradient diamond‐like carbon (FGDLC) films were fabricated using a novel pulsed laser deposition technique to enhance adhesion strength. A 355 nm picoseconds laser beam was split into two beams, and the power of each split beam was changed individually by a motorized beam attenuator as a function of time. In this way, two laser beams with customized time‐varying powers were available for ablating two different target materials. Two beams were irradiated on graphite and 316L stainless steel targets, respectively, in a vacuum chamber, and the produced dissimilar plasmas were mixed in space before they were deposited on a stainless steel 316L substrate. Using this method, they have built FGDLC films with a thickness of ∼510 nm, where the composition changes gradually from stainless steel to DLC in the direction of deposition. They have confirmed that FGDLC films show much higher adhesion strength than normal DLC films.

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In the study of Fominski et al [37], Bi‐layer W–Se–C/diamond‐like carbon (DLC) and WSex/DLC coatings were obtained by standard and shadow‐masked pulsed laser co‐deposition from WSe2 and graphite targets. W–Se–C coatings appeared as nanocomposites containing quasi‐amorphous WSe2, WC, spherical β‐W nanocrystalline particles encapsulated in WSe2 amorphous shell, and amorphous carbon phases. In WSex/DLC coatings, the formation of chemical bonds between W and C atoms was noticed at the interface. An increased of the C concentration over 40 at.% increases hardness and elasticity (up to 2 times at ~ 60 at.%C), and the Se/W ratio was always close to 1.4. The use of shadow‐masked configuration avoided the deposition of micro‐ and nanoparticles.

However, this method led to a substantial increase of the Se content (Se/W ≥ 4), and the coatings became softer.

Conclusions 

Metal implants are mainly fabricated from stainless steels, cobalt base or titanium base alloys. Metal implants must have excellent corrosion resistance as well as good biocompatibity.



Biocompatibity is achieved on a metal implant by the application of proper surface treatment. Laser surface treatment is considered as a new and promising technology for surface treatments.



Laser surface treatments performed on metal implants are mainly of two types; the first concerned with the deposition of hydroxyapatite (HA), bone‐like material while the second focused on the deposition of diamond‐ like carbon (DLC). DLC coatings are formed only under vacuum.

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