Hydroxyapatite (HA) coatings for biomaterials

5 Hydroxyapatite (HA) coatings for biomaterials P. CHOUDHURY and D. C. AGRAWAL, CSJM University, India

Abstract: Hydroxyapatite (HA) is a major constituent of hard tissues such as bone and teeth. Synthetic HA is therefore of great interest as a transplant material to replace these tissues. While the use of HA in unloaded implants and dental implants (with reinforcing metal posts) is common, the poor mechanical properties of HA do not allow its use in load-bearing implants. In such cases, a coating of HA on a metallic implant is preferred. In this chapter, methods of preparing HA coatings such as plasma spraying, sol-gel process and biomimetic growth are discussed in some detail, while other methods are briefly discussed. Key words: hydroxyapatite, coatings, plasma spraying, sol-gel, biomimetic.

5.1

Introduction

During the last few decades, materials for biomedical applications have received greater attention in the scientific community, primarily due to the fact that suitably designed biomaterials are capable of replacing, reconstructing and regenerating human/animal body tissues for long-term use, without many toxic or inflammatory effects. In specific applications (such as hard tissue replacements), materials development is aimed at maintaining a balance between the mechanical properties of the replaced tissues and the biochemical effects of the material on the tissue. For the clinical success of the materials, both areas are equally important. However, in most biological systems, a range of properties such as biological activity, mechanical strength, chemical durability, etc. is necessary. Therefore, only a designed material, with a complex combination of properties, can achieve such clinical needs. A wide variety of biomaterials have been developed, including different types of composites and coatings. The major issues in the development of biocompatible coatings are coating/substrate adhesion strength being limited by factors like non-uniform coating thickness, delamination due to mismatch of coefficient of thermal expansion (CTE) and weak interfacial bonding between substrate and the coatings (Epinette and Manley, 2004). Ceramics are used as bone filling materials and load-bearing components in various orthopedic joint replacements. Various composites – for example, 84 © Woodhead Publishing Limited, 2012

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metal–ceramic, ceramic–polymer and ceramic–ceramic – are being developed in order to achieve better chemical resistance and mechanical strength (Klein et al., 1993). Among bioceramic materials, HA, with a composition similar to bone and teeth, is one of the most biocompatible and is widely studied for various applications requiring good bioactive properties (Hench, 1998; Lacefield, 1993; Sardin et al., 1994). Although HA is a highly biocompatible and bioactive material, it suffers from poor mechanical properties, thereby limiting its load-bearing applications as a bulk monolithic material. Therefore, for all practical purposes, hydroxyapatite (HA) could be used in combination with another metal/ ceramic phase, which can improve the mechanical properties of HA without decreasing its biocompatibility. HA, or calcium phosphate (CaP)-based bioceramics in general, are popularly used as coatings on orthopedic and dental implants of metals/their alloys (Souto et al., 2003). The use of coated bioimplants combines advantageous properties of both coating and substrate materials. For example, bioactive ceramic coatings on metallic implants combine the good strength of the metal with the good bioactivity of the ceramic coating. The coating deposition route also influences the physical properties of the coatings. For example, it is reported that HA-containing glass coatings on Ti dental implants have better adhesion than flame-sprayed HA coatings. HA-containing glass coatings have increased abrasion resistance, improved aesthetics (color, etc.) and enhanced bioactivity.

5.2

Hydroxyapatite (HA) coatings

The poor mechanical properties of HA limits its use in the bulk form. This section describes the various aspects of HA coatings on metallic substrates such as, the desirable properties and preparation techniques.

5.2.1 Introduction HA in the bulk form exhibits poor mechanical properties, like low fatigue resistance, brittleness, low strength and toughness. HA has a bending strength less than 100 MPa, which is considerably lower than that of the native bone. The presence of porosity, which may be in the form of micropores (<1 µm diameter) due to incomplete sintering, or macropores (>100 µm diameter) created to allow bone growth, further affects the tensile and compressive strength and fatigue resistance (Cook et al., 1992). In physiological solutions, the Weibull modulus (m) of HA implants has a low value (m = 12) indicating a low reliability under tensile loads (Pajamaki et al., 1995). These drawbacks prevent the usage of HA in load-bearing applications. Presently, bulk HA and other CaP bioceramics are used as powders, small unloaded implants, dental implants (with reinforcing metal posts), low-loaded porous

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implants or bioactive phases in polymer-bioactive ceramic composites. Other applications include repair of bony defects, repair of periodontal defects, alveolar ridge augmentation, ear implants, eye implants, maxillofacial reconstruction, spine fusion, bone space fillers, implant coatings, adjuvant to uncoated implants, etc. (Metzger et al., 1982; Niwa and LeGeros, 2002). In the future, CaP bioceramics may be used in drug delivery system, growth factor carriers, effective carriers of bioactive peptides or bone cells, periodontal ligament cells, mesenchymal cells and hence they will be useful in tissue engineering for the regeneration of hard tissues (Kuboki et al., 1998; Ohgushi and Kaplan, 1999). The development of good mechanical strength between an implant and bone (osseointegration) decides the success or failure of the implant. Metal implants have a poor ability to bond to bone and to promote bone growth. Moreover, such implants undergo corrosion and release metal ions into the tissue (Hayashi et al., 1990). Hence, coatings of metal implants with bioactive ceramics like HA not only prevent corrosion, but also increase the growth rate of tissue within the pores (Balamurugan et al., 2002; Chai and Ben-Nissan, 1999; Lange and Donath, 1989; Liu et al., 2001). In this approach, bone ingrowth occurs in the porous implants (biological fixation) and attachment also occurs by direct chemical bonding with the bone (bioactive fixation). HA-coated implants increase the quality of adhesion of structural prostheses and hence reduce particle release from the metal, playing an important role in considerably reducing the number of rejection cases (Manso et al., 2000). The ideal HA coating for orthopedic implants is one with low porosity, strong cohesive strength, good adhesion to the substrate, a high degree of crystallinity, high chemical purity and phase stability. Most commonly, the level of crystallinity in HA coatings is about 65–70% (Tsui et al., 1998a). Highly crystalline HA coatings show low dissolution rates in vitro, with less resorption and more direct bone contact in vivo. Amorphous HA undergoes rapid dissolution in the physiological environment. Therefore, HA with low crystallinity quickly becomes weak and may promote inflammatory responses (Tsui et al., 1998b). It is therefore desirable to have a high degree of crystallinity in HA coatings, although the presence of a small amount of amorphous HA at the coating surface may promote beneficial physiological activity. Godley et al. (2004) reported that a necessary condition for a biomaterial to bond with the living bone is the formation of a biologically active bonelike apatite on its surface. They demonstrated in their work that chemical treatment can be used to create a CaP surface layer, which might provide the alkali-treated Nb metal with bone-bonding capability. A similar CaP layer will form upon implantation of alkali-treated Nb into the human body which should promote the bonding of the implant to the surrounding bone.

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5.2.2 Preparation and properties The most popular method of fabricating HA coatings is the plasma-spraying technique. Other available methods are hot isostatic pressing, spray painting, flame and oxy-fuel combustion spraying, magnetron sputtering; ion-beam deposition, chemical deposition under hydrothermal conditions, electrochemical deposition; metal-organic chemical vapour deposition (CVD), sol-gel, pulsed laser deposition; polymeric process; and electrophoresis. HA coatings have been applied to metallic materials like titanium alloys or Co–Cr–Mo alloy, as well as to carbon implants, sintered ceramics like ZrO2 and Al2O3 and even to polymers like polymethyl methacrylate (PMMA). Some of the coating techniques are summarized in Table 5.1 along with the advantages and disadvantages of each process. Table 5.1 Different coating techniques for HA Technique

Thickness

Advantages

Disadvantages

Dipcoating

0.05–0.5 mm

Requires high sintering temperatures, thermal expansion mismatch

Sputter coating

0.02–1 μm

Inexpensive coatings applied quickly, can coat complex substrates Uniform coating thickness on flat substrates

Pulsed laser deposition Hot pressing (HP) and hot isostatic pressing (HIP) Electrophoretic deposition

0.05–5 μm 0.2–2.0 mm

Produces dense coatings

Cannot coat complex substrates

0.1–2.0 mm

Uniform coating thickness, rapid deposition rates, can coat complex substrates High deposition rates

Difficult to produce crack-free coatings, requires high sintering temperatures

Thermal spraying 30–200 μm

Sol-gel

<1 μm

Line of sight technique, expensive, time consuming, cannot coat complex substrates, produces amorphous coatings As for sputter coating As for sputter coating

Can coat complex shapes, low processing temperatures, relatively cheap as coatings are very thin

Line of sight technique, high temperatures induce decomposition, rapid cooling produces amorphous coatings Some processes require controlled atmosphere, expensive raw materials

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5.3

HA coatings by plasma spraying

This section describes in detail the plasma-spray process as applied for HA coatings on different substrates, their characterization and the effect of plasma-spraying variables on microstructural features.

5.3.1 Plasma-spray process The plasma-spraying process is widely used for coating metallic substrates, usually Ti or Ti alloys, by HA (Weng et al., 1995, 1996). In this process, dense HA particles or spherical beads are fed into an electric arc-plasma gas atmosphere with temperatures of 10 000–30 000°C (Groot et al., 1987; Wen et al., 2000). The HA particles undergo partial melting during the passage through the high-temperature atmosphere and are then deposited on the substrate (Weng et al., 1995). This process leads to the formation of several phases, such as amorphous calcium phosphate (ACP), α-tricalcium phosphate (α-TCP), β-tricalcium phosphate (β-TCP), tetracalcium phosphate (TTCP) and CaO. Plasma-sprayed HA coatings on metallic substrates, thus, always consist of an inhomogeneous microstructure (Yan et al., 2003). Coating thicknesses are in the range of 30–200 µm.

5.3.2 Effect of plasma-spraying variables The effect of increasing plasma power levels has been studied by Tsui et al. (1998a). With the increase in plasma power level, the crystallinity and OH− ion content of the coating decreases, the amount of non-HA CaP compounds increases and the porosity level and extent of microcracking decreases. These changes are due to the increased degree of particle melting and subsequent freezing of molten particles to an amorphous phase. In the plasma-spray technique, tailoring the powder feedstock morphology and properties through suitable conditioning processes can help in deposition efficiency and improve the coating structure (Fu et al., 1998). The integrity, microstructure, surface profile and degree of crystallinity of the coating depend to a large extent on the morphology, structure and size of the powder (Khor and Cheang, 1994). Stability of the feedstock affects the flow behavior, deposition consistency, phase content and degree of crystallinity in the as-sprayed coating. Better flow behavior is observed for powders with a narrow particle size range compared with mixed sizes of large and small particles. Most of the HA powders are of the agglomerated, spraydried type. Particles used are usually spherical, with an average diameter of 50–60 µm. During the thermal spraying process, the very high heating rates and temperatures exceeding 10 000°C melts the majority of the powder. Large

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particles are partially melted while small particles may have a layer of vaporized material adherent to the liquid droplets. The heated powder impacts the substrate at a high velocity and cooling rate (about 106 Ks−1), resulting in the formation of the coating. The solidification process of the molten droplet is, therefore, a non-equilibrium phenomenon. The combination of these conditions leads to the formation of metastable phases. The resultant microstructure consists of macro- and micro-porosity, fused and partially molten particles with limited interparticle cohesion (Wen et al., 2000). Porous particles form hollow spheres resulting in increased coating porosity. Wang et al. (1999) studied the effect of particle size on the plasma-sprayed coatings. HA powders produced by the precipitation method were spraydried at 200°C, spheroidized by combustion flame spraying into distilled water and, finally, oven dried for plasma spraying. Functionally gradient coatings using (i) spheroidized HA (SHA) powders of different particle sizes, (ii) mixtures of SHA and spherical α-TCP powders and (iii) mixtures of SHA and TiO2 powders were prepared. The as-sprayed SHA coatings consisted of HA and non-HA phases like α-TCP, CaO and TTCP resulting from the thermal decomposition occurring in the plasma flame. A dense coating was obtained with small particle sizes. The microstructure of the as-sprayed SHA coating consisted of several randomly stacked lamellae. Microcracks and micropores were observed at the intersection of these lamellae. Tong et al. (1996) observed that the amount of additional (non-HA) phases were present in higher quantities in the as-sprayed coatings formed with smaller particles. Crystallinity of the coatings decreased with particle size in the small and large size range. The surface of the coatings made with 150 µm particles was covered with unmelted particles but these were absent for the coatings made with 180 µm particles. Thus, the latter had a higher melt ratio. The deeper penetration and a higher residence time of these large-sized particles in the plasma flame resulted in a higher melt ratio. Large particle sizes result in higher porosities, larger unmelted cores, less amorphous content and additional phases. Smaller-size starting powders are sufficiently melted resulting in lower porosities, relatively smaller unmelted cores, more amorphous material and additional phases. Three powders, namely, calcined HA (CHA), spray-dried HA (SDHA) and SHA were used by Cheang and Khor (1996). Particles in the size range of 53–75 µm were used for plasma spraying. The coating microstructure produced using CHA powders consisted of micro- and macropores, irregular splats and unmelted particles with limited interlayer adhesion. Limited interparticle cohesion led to poor structural integrity in these coatings. ACP, TTCP, TCP and calcium oxide were the phases present in addition to crystalline HA. The use of SDHA powders led to a reduction in overall porosity owing to better flow behavior of the spherical powder. The impact of

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irregularly sized particles led to phase inconsistency. The SDHA coatings had a lower crystallinity of HA along with lesser amounts of calcium oxide. The use of SHA powders resulted in a regular formation of neatly stacked disc-like splats producing a smooth and flat surface profile and the lowest amount of calcium oxide. The cross-sectional view of the coating microstructure consisted of a lamellar structure without macrovoids. The coating integrity was significantly improved by the presence of good interlamellar contact and the absence of macropores and unmelted particles. In the coating formed by Fu et al. (1998), the first layer consisted of particles in the size range of 20–45 µm to provide adhesive strength. The top layer was sprayed with HA particles in the size range of 75–125 µm to provide biocompatibility, while the middle layer was composed of powder in the size range of 45–75 µm. The surface of the as-sprayed coating consisted of undulating surfaces of unmelted, partially melted and fully melted splats. A number of microcracks and micropores resulting from the spray process and thermal residual stress were also visible on the surface. A ceramic slurry mixing process has been used to prepare composite powders of HA/Ti6Al4V for plasma-spray coatings on Ti6Al4V substrates (Gu et al., 2003). The resulting composite coatings were found to be denser with a higher bond strength than pure HA coatings. The morphology of the as-received composite coatings consisted of a rough, heterogeneous and melt-like structure. Some microcracks were observed on the surface which formed both due to a thermal expansion mismatch between the coating and substrate and the release of thermal stresses generated during cooling.

5.3.3 Microstructure of plasma-sprayed coatings As mentioned earlier, the microstructure of the plasma-sprayed coatings is quite inhomogeneous. Wen et al. (2000) have observed two metallographically different regions in HA coatings on Ti6Al4V: a ribbon-like region surrounded by a relatively smooth area. The ribbon-like region had a granular surface appearance. Secondary ion mass spectrometry did not reveal any significant difference in chemical composition of the two regions. Analysis based on nano-indentation and micro Raman techniques revealed the existence of an amorphous phase surrounding the ribbon-like crystalline areas. This amorphous phase was formed due to the rapid cooling of the molten HA phase. The use of micro Raman spectroscopy proved the existence of a chemical gradient in the coating thickness direction. The characteristic bands of OH− and PO43− indicated dominance of the ACP at the coating/Ti interface. It thus indicated that crystallinity decreases from the surface to the interface in the as-sprayed coatings (Abdel-Aal et al., 2008; Qiyi et al., 2003).

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Yang et al. (2009) coated HA (45–60 µm particles) on laser gas-nitrided pure titanium and grit-blasted pure titanium substrates using a plasma spraying technique. Microstructures of the coating were mainly composed of HA, ACP and some minute phases of TCP, TTCP and calcium oxide. For the long-term use of HA coatings, where highly crystalline HA phase is necessary, post-deposition treatments such as sintering or hydrothermal treatment are used to restore biocompatibility. The effect of heat treatment on the composition, crystallinity and other properties of the plasma-sprayed coatings has been studied by Qiyi et al. (2003). The as-received coatings consisted of HA as well as other CaP phases (Gross et al., 1998a; Kim et al., 2004). The TCP and TTCP did not undergo any change during heat treatment at high temperature (650°C, 900°C) in a vacuum (Knowles et al., 1996). Only the amorphous HA changed to a crystalline form. However, if water vapor is present during heat treatment then these phases as well as the amorphous phase tend to change to crystalline HA. A treatment at 120°C in water vapor yielded a pure HA phase along with a significant increase in the crystallite size. The amorphous phase in an HA coating deposited on grit-blasted Ti by plasma spraying was studied by Gross et al. (1998b). A dehydroxylated CaP was the major phase of the amorphous HA. On heating, the hydroxyl-rich areas crystallized to form HA. Diffusion of hydroxyl ions from the surroundings further increased the amount of the crystalline phase. Crystallization occurred over a range of temperatures depending on the hydroxyl content of the amorphous phase and the partial watervapor pressure. The activation energies for the crystallization to HA and oxyapatite are 274 and 440 kJ/mol, respectively, while that for diffusion of hydroxyl ions is 230 kJ/mol. The cell response to the coating seems to be less sensitive to the microstructure and the amount of crystallinity than to the surface roughness. Kim et al. (2004) investigated the effect of surface modification of the plasmasprayed HA coatings on the attachment and proliferation of the bone cells. It was found that the effect of modification of the coating by using different plasma spray parameters could be correlated to the surface roughness rather than to any changes in the phases and microstructure. Chou and Cheang (1999) characterized the microstructure of plasmasprayed HA-10 wt% ZrO2 composite coatings on titanium. The ZrO2 used was stabilized with 8 mol% Y2O3. Transmission electron microscope (TEM) studies revealed the presence of phases such as HA, ACP, α-TCP, ZrO2 and minor transformed CaZrO3. The cubic ZrO2 phase is maintained during plasma spraying and the ZrO2 particle bonds well to the CaP matrix with a local crystallographic relationship. The composite coatings showed the presence of more unmelted small particles. CaP and ZrO2 react rapidly to form CaZrO3.

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5.4

Properties of plasma-sprayed coatings

This section reviews the important properties of HA coatings produced by plasma-spray process, such as mechanical properties and response to liquid media.

5.4.1 Mechanical properties Several researchers have studied the effect of coating quality on the mechanical properties of the plasma-sprayed coatings. The stability of the implant/coating and the coating/bone interfaces strongly affect the performance of the implant. The implant/coating bond is largely physical, while the coating/bone bond is physicochemical in nature. The performance of both interfaces is related to the coating properties. Microhardness, Young’s modulus, fracture toughness and bond strength are the mechanical properties most commonly evaluated. There is a lack of data on the Young’s modulus of sprayed HA coating. The absolute values of the Young’s modulus determined by Tsui et al. (1998b) were less than 6 GPa, which is possibly due to the presence of extensive microcracking and a high level of porosity. The Young’s modulus was found to be considerably lower in tension than in compression, which is attributed to the opening up of microcracks under tensile loading. Significant improvement in Young’s modulus (~55 MPa) and tensile adhesion strength (~28 MPa) was observed in HA/Ti6Al4V composite coatings by Dong et al. (2003). Residual stress levels were predicted to be ~20–40 MPa tensile at room temperature (Tsui et al., 1998b). This relatively low value is attributed to the low coating stiffness. Owing to the low interfacial toughness (common in these systems), the energy released during stress relaxation may be sufficient to cause interfacial debonding, particularly for thick coatings. Hence, it may be useful to minimize the substrate temperature during spraying to further reduce the residual stress level. Particle size, morphology and composition play significant roles in the mechanical behavior of the plasma-sprayed coatings. In the study by Wang et al. (1999), the average microhardness values were higher for relatively smaller particle sizes. The addition of TiO2 and α-TCP significantly increased the fracture toughness and microhardness of HA coatings, with TiO2 offering greater toughening efficiency. Thus, methods like the introduction of a toughening phase can be used for effectively improving the performance of HA coatings. A relatively low microhardness value was observed in the transitional area between adjacent layers prepared using different sizes of powders. This is possibly due to the relatively high residual (tensile) stress at these locations. With an increase in the feedstock particle size, fracture toughness of the coating decreases. The use of spheroidized powders increases the density and, hence, the mechanical properties of the coatings.

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The effect of the thickness of plasma-coated HA on the mechanical fatigue behavior of the HA/Ti system has been evaluated by Lynn and DuQuesnay (2002). The coatings revealed 5% porosity by volume and cylindrical pore sizes in the range of 0.1–10 µm. It has been found that a coating thickness of 0–100 µm does not affect the fatigue life behavior of Ti6Al4V substrates. 150 µm thick coatings reduce the fatigue resistance of similar substrates. 25–50 µm thick HA coatings do not show any observable delamination during fatigue or final fracture, while 75–150 µm thick coatings spall (i.e., flakes removed from the surface) following the initiation of the first fatigue crack in the substrate. For all coating thicknesses, the critical fatigue crack initiates at the edge of the substrate. The extent of damage caused to the substrate by the larger cracks in the thicker coatings is expected to be greater than that caused by the smaller cracks in the thinner coatings. Reduction in fatigue life of the thicker coatings is due to stress relief in the near-surface regions of the substrate due to increased thermal energy inputs. The higher heat inputs result from accelerated conductive and radiative heat transfer mechanisms during the application of such thick coatings, and from the extended times required to deposit them. HA coatings are subjected to wear in several applications, like pedicle screws in spinal implants, external-fixation pins, shoulder and hip prostheses. The greatest degree of abrasion is in the femoral stem used as part of hip prosthesis. Different coating characteristics were obtained by using two different powders and by changing the spray distance (Gross and Babovic, 2002). Three types of coatings were prepared on stainless steel plates dense and smooth, slightly porous and rougher, and slightly porous and rough. The abrasion resistance of the thermally sprayed HA coatings, about 100 µm thick, was investigated using a pin-on-disk arrangement in unlubricated conditions. The pin was made of wood, a bone analogue, in terms of hardness and elastic modulus. All the coatings, irrespective of roughness, exhibited a weight loss along with a decrease in surface roughness during abrasion. A larger weight loss occurred for higher surface roughness and the use of heavier loads. Particulates broke off from the elevated areas of the coating during abrasion. Abrasion and wash-out are responsible for particle formation during initial stages of implantation. During insertion of the coated implant, abrasion against bone will produce particulates. The surrounding physiological solution dissolves materials from the highly stressed areas, thus, weakening the raised areas. Wash-out results from the preferred dissolution of ACP and, thus, depends on the phase composition of the outer layer of the coating. On the other hand, abrasive release of particles will depend on the type and number of raised areas. The coating roughness, traverse distance of the coating relative to the bone, the insertion force and the residual stress within the coating determine the release of particles. Though coating roughness helps in implant fixation, it also leads to particle production on the coating surface.

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The appearance of shearing micromovements at the interface between the implant and the bone due to large differences in elastic modulus of the two materials in contact induce fretting wear and sometimes, fatigue cracks (Yang et al., 2009). This process causes the early failure of joint prosthesis. Fu et al. (1998) carried out tests for fretting wear on 200+/−20 µm thick HA coating on Ti6Al4V using a ball-on-flat fretting apparatus. The material to be investigated was the flat specimen, whereas the upper specimen consisted of a hardened steel ball. Tests were conducted at room temperature (25°C) and at 80% relative humidity. The frequency of tests was 5 Hz at loads of 5 and 10 N. The amplitudes of oscillation were 50, 100 and 200 µm, respectively. The coefficient of friction for the HA-coated alloys was 0.7–0.8. HA coatings underwent fretting wear under dry sliding conditions mainly by delamination and abrasive wear. At less than 5 × 106 fretting cycles, the difference in surface roughness and wear mechanism led to a little higher fretting wear resistance of the HA coatings than the titanium substrate. As fretting cycles increased beyond 5 × 106, severe wear of HA coatings occurred by crack initiation from pores inside the coating and increased ploughing by debris. This resulted in a significantly lower fretting wear resistance of the HA coating as compared with the titanium substrate. In the investigation of Yang et al. (2009) reported earlier, the threedimensional (3-D) TiN dendritic scaffold structure produced on the pure titanium surface, using a laser gas-nitriding technique followed by acid etching could help in the improvement in interfacial adherence significantly as compared with that on the grit-blasted surface. The bonding strength on the etched laser gas nitride substrate was 1.8 times higher than that on a gritblasted substrate. The results reveal that the measured bonding strength was a combination of adhesive and cohesive strength. As all the coatings were obtained by a plasma-spraying technique using the same spraying parameters, the structure of the coatings was similar and hence cohesive strength was the same for all the coatings. Thus, the apparent differences in the bonding strength resulted from the differences in adhesive strength. The improvement in adhesive strength resulted from a locking mechanism facilitated by the extra surface area of the 3-D dendritic network. Dong et al. (2003) investigated plasma-sprayed HA/ Ti6Al4V composite coatings containing ~48 wt% HA. Such coatings demonstrated attractive tensile adhesion strength (~28 MPa) and improved Young’s modulus (~55 MPa). Experimental results showed ACP and fine HA grain formation during rapid solidification in the as-sprayed composite coatings. The amorphous Ca3(PO4)2 underwent significant but incomplete crystallization after heat treatment of the coatings at 600°C for 6 h. After immersion in SBF (simulated body fluid) for two and ten weeks, observations confirmed that interfaces inside the coating maintained good microstructural integrity.

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5.4.2 Response to liquid media Dissolution properties of the CaPs differ due to differences in composition, crystal size and sintering temperature (Tong et al., 1995). HA has greater stability than other CaPs at higher pH. β-TCP is more stable at pH > 6.36. Ca(OH)2 has a tendency to form at a pH greater than 8.0 (Vatanatham and Kimura, 2001). Dissolution properties of bicalcium phosphate depend on the HA/β-TCP ratios: the higher the ratio, the lower the dissolution rate. Calcium and phosphorus ions redistributed during the dissolution of HA may be used by the surrounding tissues in the manufacture of new natural bone (Gledhill et al., 2001). The morphologies and the different phosphate phases present in HA also determine its solubility in vivo. Highly amorphous coatings are expected to dissolve most rapidly. The rapid dissolution results in structural decohesion within the coating, resulting in a very short service life. A semi-crystalline system is more effective in which the amorphous regions undergo rapid dissolution releasing bone-forming ions, while the crystalline phase provides the structural integrity (Kim et al., 2003). Several investigators have studied the changes in plasma spray coated samples immersed in various media like water, SBF, Hanks’ balanced salt solution (HBSS), Ringer solution, etc. In an in vitro study of HA/Ti6Al4V composite coatings in SBF solutions (Gua et al., 2003), the coatings were found to undergo two biointegration processes – that is, dissolution during the initial four weeks soaking in SBF and the subsequent bone-like apatite crystal precipitation. The composite coatings showed better long-term mechanical stability than the pure HA coatings in a physiological environment. Ha et al. (1998) investigated the response of vacuum plasma-sprayed HA coatings upon exposure to a simulated physiological environment. The immersion periods varied up to 28 days. HA was the main component in the as-received coatings along with some amount of β-TCP. The surface morphology consisted of a rough, heterogeneous, melt-like structure. After immersion for one day, the morphology changed to a needle-like structure. Measurement of the concentration of the immersion solution showed that dissolution of TCP starts on the first day. At the same time, carbonate containing partially crystalline CaP was observed to precipitate. The precipitation of this new CaP layer was slowed down by the presence of proteins which adsorbed on the coating and occupied the probable sites of precipitation. The integrity of the HA coating was not affected adversely which showed high stability of the HA coatings produced by vacuum plasma spraying. HA-coated samples prepared using different-sized starting powders were immersed in deionized water at room temperature for one month

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(Liu et al., 2002). Severe detachment and degradation occurred in the coatings with the highest crystallinity accompanied by a low calcium loss. A uniform degradation and no detachment were observed in the coating made from the smaller-sized particles. However, the higher amount of the amorphous phase in these coatings resulted in greater amount of the calcium release. The two possible methods of degradation of coatings are dissolution of the soluble phases like α-TCP and amorphous phases present in coatings and the other is the loosening or detachment of unmelted cores of starting powder in the amorphous matrix. When the amorphous matrix dissolves, the more crystalline fragments of the coatings undergo detachment. This causes more rapid degradation than that resulting from CaP dissolution alone. On the other hand, mechanical detachment or loosening of the unmelted cores (crystalline HA) hardly increases the calcium in solution. Hence, scanning electron microscope (SEM) observation showed severe degradation and detachment in coatings with large particle sizes while the calcium content in solution is lower in these cases as compared with the coating with smaller particle sizes. Thus, microstructure plays a significant role in determining the stability of coatings. Coated specimens were aged at room temperature for 30 min or 3 h in distilled water and 0.2 M sodium phosphate buffer (pH = 7.2; Drummond et al., 1998). Crystal growth occurs on the HA-coated surface just after 30 min of incubation in the sodium phosphate buffer solution. The crystals are formed when free calcium of the HA coatings comes in contact with the phosphate ions of the buffer solution. Little difference was observed in the surface morphology after aging for 3 h in distilled water. The bulk analysis did not show any difference in either of the solutions with aging times. The difference in observation from that of Liu et al. (2002) could be due to the use of a different particle size or shorter times used here. In static SBF (Im et al., 2007), the coating first dissolved in the solution. A high concentration of Ca and P in the medium resulted in the growth of crystals without any preferential orientation. In the flowing system, the bonelike apatite layers appeared radially needle-like for the samples treated at 120°C in water vapor, and fiber-like nets and particles for those treated at 650°C in a vacuum. Here, the SBF was refreshed periodically, resulting in a lower concentration of Ca and P ions in the solution. The lower ion concentration restricted the crystal growth and also led to preferential growth of crystals. Yang et al. (2009) immersed HA-coated titanium in Hank’s solution for 20 days under static conditions and at 37°C. The immersion in the solution caused a slight decrease in the bonding strength between the coating and substrate. Qiyi et al. (2003) studied the effect of soaking the plasma-sprayed coatings in Tris-HCl buffer (pH 7.4) and SBF. The SBF was refreshed after each measurement but the Tris-HCl buffer was not replaced during

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the entire soaking period. The coatings with smaller crystallite sizes were found to undergo more rapid dissolution indicating that the attack primarily takes place at the grain boundaries. However, the imperfections created due to the rapid cooling rate may also contribute to the rapid dissolution. The presence of untransformed phases in the vacuum-treated samples led to their higher solubility than the vapor-treated coatings. The rough surface after dissolution makes the nucleation of new phases from solution easier. Consequently, due to the large number of nuclei on the former, the final grain size was smaller. The dissolution behavior of plasma spray HA-coated Ti6Al4V substrates in HBSS at 37 +/−1°C was studied by Sousa and Barbosa (1996). A comparison was made between different surface conditions: (i) a polished and passivated surface, (ii) grit-blasted and (iii) grit-blasted with HA coatings (50 and 200 µm). A passivated polished Ti surface proved to be superior to a 50 µm HA coating on the grit-blasted surface. However, increasing the thickness of the coating to 200 µm resulted in better protection than the passivated polished surface. Crevice attack was not induced at the metal/HA interface. After immersing the coated samples for six months in HBSS at 37°C, partial detachment of the HA coating on grit-blasted surfaces occurred for both the thicknesses. One of the possible mechanisms proposed for the detachment of the coatings is the hydrolysis of metal ions in solution with a consequent decrease in pH. This acidification can promote rapid dissolution of the HA at the interface leading to its detachment. The other explanation was based on the passive film (TiO2) formation reaction that reduces the pH at the Ti/ HA interface. Owing to the lack of a strong chemical bond between HA and titanium, physical bonding is important. Physical bonding requires the use of rough surfaces, which in turn increases the risk of metal dissolution. The smaller thickness of 50 µm was not sufficient to counterbalance the effect of increased surface roughness. Chemical analyses of the solutions showed that the amount of metal ions in the solution was below the detection limit of the instrument, indicating that they might have been entrapped within the HA. Metal phosphate formation or incorporation of metal ions in the HA structure prevents the release of metallic species into the HBSS. The bond strength of the composite coating was much higher than that of the pure HA coatings. The bond strength in the HA/ Ti6Al4V composite coating exposed to SBF at 37°C (Gu et al., 2003) was much higher than that of the pure HA coatings. The higher bond strength in the composite coating resulted from metallic joints among Ti6Al4V lamellae within the coating as well as the strong bonding of the Ti6Al4V in the incipient layer of splats onto the grit-blasted substrate of the Ti alloy. During soaking, tiny granular precipitates grew on the surface. Impurity phases like CaO were completely dissolved during soaking. The calcium concentration in the immersion solution initially increases due to dissolution followed by

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an exponential decrease indicating continuous precipitation (bone-like apatite). More cracks and micropores were observed in the coating after soaking in SBF, as compared with the as-prepared coating. The apatite layer formed on the coatings can be useful in bonding with the living tissues and thereby increasing the longevity of the coatings during implantation in vivo. No additional phases were formed and the crystallinity of HA increases with increasing duration of soaking in SBF. Pure HA coatings had similar trends of dissolution as that of the composite coating, but the rate of dissolution of the CaO is higher. Factors like phase purity, crystallinity, porosity and thickness of material affected the dissolution of ions in the composite coating. Dissolution resulted in a decrease in bond strength, microhardness and Young’s modulus of the coatings with increasing soaking time. Chemical dissolution of the coatings weakens the bonding of the lamellae in the coating and bonding at the coating/substrate interface. Slower dissolution of the composite coating results in significantly lower decrease in bond strength. The superior mechanical stability of the composite coatings as compared with pure HA coatings showed their better long-term stability in a physiological environment.

5.5

Biomimetic HA coatings

The synthesis of HA coatings by a biomimetic process at a physiological temperature of 37°C has become a very promising approach. Such films can be prepared by soaking the substrate (Cho et al., 1995; Hata et al., 1995; Pereira et al., 1995; Tanahashi et al., 1994, 1995, 1996) (silica gel, Ti, alumina and polymers) in SBF. Alternatively, a plate of CaO–SiO2 glass can be placed close to the Ti substrate during soaking in SBF. The glass acts as a nucleating agent to form an HA coating on an untreated Ti substrate. The HA coating formed by a biomimetic process may not have good adhesion strength (usually about 10 MPa). Surface treatment plays an important role on the biomimetically deposited HA coating as described in this section. Table 5.2 summarizes the coating characteristics obtained by different investigators using different substrate pretreatments. Zhang et al. (2003) used porous titanium substrates in their study. The samples, cleaned in acetone, alcohol and boiling deionized water, were further immersed in a mixture of 18% HCl and 48% H2SO4 for 1 h and in NaOH at 70°C for 4 h. The thoroughly rinsed samples were soaked at 4°C for 24 h, in a solution prepared by dissolving HA powder in 0.1 mol/l HCl with the pH being adjusted to 7.3 by adding tris-hydroxymethylaminomethane (TRIS) buffer. Subsequent to the deposition of crystal seeds of HA on the substrates, the solution temperature was increased to 37°C and held for 24 h to obtain uniform coatings. Coatings on porous titanium having uniform macropores (about 200 µm interconnected completely)

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Table 5.2 Summary of the coating characteristics obtained by different investigators using different substrate pretreatment methods (biomimetic)

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No.

Substrate treatment

1

Porous Ti cleaned in acetone + HA powder dissolved in HCl, 50% porosity, macropore sizes pH adjusted to 7.3 with TRIS alcohol + deionized water; 20–800 μm, spindle-shaped crystal buffer treated in HCl+H2SO4 for 1 h 100 nm dia and 500 nm long and NaOH for 4 h Ti cleaned in ethanol, dried Modified SBF at pH 7.3 Inhomogeneous and non-uniform at 100°C. Some samples adjusted with TRIS buffer layer of HCA after 20 days alkali-treated, others acidexposure etched followed by alkali treatment; all samples washed and dried at 100°C Alkali pretreated Ti SBF with ionic concentration Flake-like (carbonate containing 1.5 times human blood HAp) porous apatite consisting of plasma at 36.5°C very fine nanoapatite crystallites on alkali-treated surfaces only. Coating thickness increased with soaking time Ultrasonically cleaned Ti acid Static SBF at 37°C for 10 days Rate of apatite formation not affected pretreated followed by max by lower content of Na+ in surface layer alkali treatment at 60°C for 24 h Ultrasonically cleaned Ti SBF solution at 36.5°C and pH 6% H2O2 solution treatment at a pH of 4–4.6 at 60°C for 3–6 h followed by treated with H2O2 solution 7.4 for 3 days of different concentrations heat treatment provides optimum for different lengths of time in vitro forming ability at 60°C

2

3

4

5

Solution used for immersion

Coating characteristics

Reference Zhang et al. (2003)

Habibovic et al. (2002)

Habibovic et al. (2002) and Ma et al. (2003)

Jonasova et al. (2002)

Shibata et al. (2003)

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Table 5.3 Ionic concentration (mM) of SBF Ions

Na+

K+

Concentration (mM)

142.0 5.0

Ca2+

Mg2+

HCO3− Cl−

2.5

1.5

4.2

HPO42−

148.8 1.0

SO42− 0.5

Source: Habibovic et al. (2002).

consisted of uniformly dispersed and cross-linked crystal flakes. The coatings on titanium having 50% porosity with macropore sizes ranging from 20 to 800 µm consisted of spindle-shaped crystal about 100 nm in diameter and 500 nm in length near the edge of the substrate. The crystals in such places were clustered. HA coatings have been formed by a biomimetic process on chemically treated Ti substrates by Habibovic et al. (2002). Some of the cleaned samples were treated in alkali while others were acid-etched followed by alkali treatment at 60°C for 24 h, washed, dried at 100°C and then soaked in modified SBF (composition given in Table 5.3) at pH 7.3 for different time periods. The growth of bacteria in SBF was inhibited by the addition of NaN3. Isolated spheroid particles (typical for apatite crystallized from SBF) were deposited on the sample surface after ten days. The increase in sample weight during gravimetric analysis indicated the formation of a surface layer which was confirmed to be hydroxycarbonated apatite (HCA). The HCA layer remained inhomogeneous and non-uniform even after 20 days in SBF. The passive oxide layer on titanium partially dissolves in NaOH to form HTiO3− while titanium is hydrated to form HTiO3−.nH2O. The negatively charged groups react with Na+ ions to produce an alkali titanate hydrogel surface layer on the titanium surface. Alkali treatment in NaOH led to the formation of a compact and a homogeneous layer, the thickness of which increased continuously with time. A uniform micro-roughened surface obtained after acid etching of Ti in HCl under an inert atmosphere provides an improved condition for in situ HCA formation. During acid etching in HCl, titanium reacts with HCl to form TiCl3 and H2. Subsequently, TiH2 is formed by the reaction between titanium and hydrogen. During exposure of acid-etched titanium in NaOH, the passive TiO2 layer (formed in atmospheric contact) dissolves to form an amorphous titania layer containing Na+ ions: TiO2 + NaOH → HTiO3− + Na+.

[5.1]

Soon after immersion in SBF, Na+ ions from the amorphous titania layer are exchanged by H3O+ ions from the surrounding fluid resulting in Ti-OH layer formation. At the same time, with increasing pH and, hence, increasing

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supersaturation of the solution with respect to apatite, the rate of apatite nucleation also increases. Thus, the two-step chemical treatment of Ti with HCl followed by NaOH provides the Ti surface with bone-bonding ability. Kokubo et al. (1996) and Ma et al. (2003) have also used alkali pretreated titanium for biomimetic coating with HA. After the alkali treatment, substrates were soaked in SBF having an ionic concentration 1.5 times higher than that of human blood plasma at a temperature of 36.5°C. Flake-like porous apatite consisting of very fine nanoapatite crystallites (60 nm) was observed on the alkali-treated surfaces but not on the untreated surfaces. The amount of HA and the thickness of the coating increased with soaking time. The coating thickness reached a saturation limit after ten days of immersion caused by the increasing distance from the growth-promoting amorphous alkali-treated layer. The alkali treatment increases the ionic activity of apatite in SBF, thus increasing the rate of nucleation of apatite. The nuclei of apatite formed on the metal surface, grew spontaneously by consuming the calcium and phosphate ions from the SBF. The apatite layer was integrated with the metal substrate through the titanate layer resulting in a strong bonding of the same, giving a shear strength of 9.5 MPa. The influence of a lower amount of Na+ in the surface layer of alkalitreated titanium on the rate of HA formation in SBF solution has been studied by Jonasova et al. (2002). Samples of ultrasonically cleaned and dried commercial purity titanium were subjected to acid (HCl) pretreatment in an inert atmosphere before the alkali treatment at 60°C for 24 h. The alkali-treated samples were washed — one to five times for 1 min in 100 mL distilled water to obtain a different surface content of alkali ions. The washed and dried samples were exposed to static SBF solutions at 37°C for a maximum period of ten days. It was found that the rate of apatite formation was not significantly affected by the lower content of Na+ in the surface layer. Hence, the use of Ti with the lowest amount of Na+ in the surface layer will release the minimum amount of alkali ions into the surrounding tissue and, therefore, will be most suitable for human body implantation. Shibata et al. (2003) have used the biomimetic approach to coat ultrasonically cleaned commercial purity titanium substrates treated with H2O2 solution of different concentrations (1, 3, 6 and 10 mass%) for different lengths of time (1, 3, 6 and 12 h) at 60°C followed by heating at 400°C for 1 h in air in an electrical furnace. The treated and washed substrates were soaked in SBF solution for three days at 36.5°C and pH 7.4. In H2O2, titania gel formation and Ti(IV) dissolution occurred competitively. The Ti–O–O–Ti and TiOOH groups in the titania gel obtained during H2O2 treatment were eliminated by thermal treatment at 400°C. The thermal treatment promoted the formation of TiO2 resulting in an increase in apatite deposition. Substrates treated with 6% H2O2 solution at a pH range of 4–4.6 at 60°C for 3–6 h and

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subsequently heat treated at 400°C for 1 h provided the optimal in vitro apatite forming ability.

5.6

HA coatings by sol-gel deposition

This section describes the different aspects of sol-gel coatings in general, characteristics of sol-gel deposited HA coatings and processing of composite HA coatings by sol-gel route.

5.6.1 Introduction The sol-gel deposition technique is a very effective coating method, especially for the deposition of thin coatings. The term ‘sol-gel’ is presently used to describe any chemical process capable of producing ceramic oxides, nonoxides and mixed oxides from solutions (Balamurugan et al., 2002; Chai and Ben-Nissan, 1999). It has been commonly used to produce glasses and oxides, but in recent times more complex materials as well as non-oxide ceramics have been prepared (Chai and Ben-Nissan, 1999). The advantages of the sol-gel process are: increased physical and chemical homogeneity due to mixing on the molecular scale, reduced firing temperatures due to small particle sizes with large surface areas, ability to produce uniform fine-grained structures, the use of different chemical routes (alkoxide or aqueous based) and their ease of application to complex shapes with a range of coating techniques – for example, dip, spin and spray coating. Use of low processing temperatures minimizes the degradation of substrate metals due to thermally induced phase transformation, microstructure modification or oxidation. Such coatings have demonstrated better structural integrity, purity and phase composition than the coatings prepared using the conventional methods, like thermal spraying. In this method, a solution is first prepared from alkoxides, metal salts or other suitable precursors. Hydrolysis and condensation reactions result in the production of nanometer-sized building blocks, which may be complexes or oligomers, and the suspension is now referred to as a ‘sol’. The coating is produced by depositing this sol using dip-, spin- or spray-coating techniques. Figures 5.1 and 5.2 are schematic illustrations of the sol-gel coating formation and the spin- and dip-coating methods, respectively. Hydrolysis and condensation reactions are accelerated in the film resulting in the formation of a 3-D gel network, which on heating is converted to the oxide under oxidizing conditions. The schematic representation for the sol-gel coating process is shown in Fig 5.3. Each coating step produces a film usually less than 0.1 µm thick; thicker coatings can be obtained by repeating the coating steps or by introducing some polymeric additives in the sol. The thin coatings are able to withstand

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Metal alkoxide solution Hydrolysis polymerization Xerogel film Coating

Heat

Dense film

5.1 Schematic illustration of sol-gel deposition technique. A solution prepared from metal alkoxides undergoes hydrolysis/polymerization reaction resulting in a suspension of oligomers/particles (sol). Evaporation of the solvent results in xerogel film, compacted into a dense film after heat treatment.

(a)

Stage 1

Stage 2

Stage 3 and 4

(b)

Dipping

Wet layer formation

Solvent evaporation

5.2 Schematic illustration of (a) spin coating and (b) dip-coating processes of the sol-gel coating technique. In the spin-coating method, a drop of the sol is placed on the substrate which spreads on the surface by its rotational motion, followed by thinning of coating by solvent evaporation. In the dip-coating method, substrate is dipped into the sol; during withdrawal, a wet layer forms on the substrate followed by solvent evaporation and subsequent thin film formation.

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Nanomedicine Calcium precursor in methanol

Phosphorus precursor in methanol

Mixing

Spin / dip coating

Prefiring

Heat treatment

Analysis

5.3 Schematic diagram of sol-gel process for preparation of HA films.

the thermal expansion mismatch, unlike thicker coatings such as those prepared by dip-coating from a powder suspension. Both inorganic and organic precursors (alkoxides) have been used to produce HA coatings. The most commonly used inorganic calcium precursor is calcium nitrate while the common inorganic phosphorus precursors are ammonium dihydrogen phosphate, phosphorus pentoxide and phosphoric acid. The commonly used organic calcium and phosphorus precursors are calcium diethoxide and triethyl phosphite, respectively. The combinations of precursors, solvents, substrates and other conditions used by various researchers are summarized in Table 5.4. In the following, we give some details of the sol-gel film preparation and properties in a few cases.

5.6.2 Characteristics of sol-gel deposited HA coatings From the summary listed in Table 5.4, it is observed that the minimum temperature for the formation of crystalline HA is 500°C if an organic solvent is used and 350°C if water is used as the solvent. However, Lopatin et al. (1998) have reported the formation of HA at 350°C even with an organic solvent. The sol must be aged for at least 24 h for the formation of monophasic HA;

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Table 5.4 Summary of the combination of variables used by various researchers for sol-gel coating of HA

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No

Precursors

Substrate

1

Calcium chloride and potassium dihydrogenphosphate

Oxidized titanium

2

Calcium nitrate and ammonium dihydrogenphosphate in ethanol Calcium nitrate and ammonium dihydrogenphosphate in ethanol Calcium nitrate and phosphorus pentoxide in ethanol

Calcium nitrate and phosphorus pentoxide in ethanol

3

4

5

Heat treatment

Other conditions

Comments

References

HA structures formed with rose-like morphology

Shibata et al. (2003)

Ti6Al4V

400–11 000°C

0.01M KOH to maintain pH = 6.5– 7.5 substrates kept in aging solution for 0.5–10 days Basic solution

Ti6Al4V

500°C

Ti6Al4V

Prefiring – 500°C Final heating – 750°C

Ti6Al4V

500°C, 750°C

Crystallinity of HA Haddow et al. improves with (1996) increasing temperature Irregular surface, 60% density

Ten coatings, some coated samples treated in SBF for 1 month

Dense, smooth morphology with spherical grains; rough, porous morphology after treatment in SBF; adhesion strength of 14 MPa Dense morphology, adhesive strength of 10 MPa, better crystallinity and adhesive strength at 750°C but porous morphology

Kim et al. (2004) and Gan et al. (2005) Qiyi et al. (2003)

Weng and Baptista (1998)

(Continued)

Table 5.4 Continued

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No

Precursors

Substrate

Heat treatment

Other conditions

6

Calcium acetate with phosphoric acid, phosphorus pentoxide and triethylphosphite

400–11 000°C

7

Calcium glycoxide and Alumina phosphorus pentoxide in glycol, n-butanol and ethanol

750°C and 1000°C

8

Calcium nitrate and triethylphosphate in 2-methoxy ethanol

Ti6Al4V

400°C; final heat Multiple-layer treatment at coatings 600°C

9

Calcium diethoxide and triethylphosphate in ethanol and ethanediol

Single crystal magnesia

10

Calcium nitrate and triehtylphosphate in ethanol and distilled water

Prefiring at 500°C, final firing at 1000°C 350°C

Acetic acid used for stabilizing Ca-glycoxide, Multiple-layer coatings

Comments

References

Temperature >600°C required for HA formation, triethylphosphite shows best wetting property Rough coatings with 10 MPa adhesion strength; length and number of microcracks increase with increasing heat treatment temperature Porous coating, formation of pure HA, first layer adhesive strength of 90 MPa Aging >24 h needed for formation of monophasic HA

Habibovic et al. (2002)

Hirai et al. (2004)

Weng and Baptista (1998) Chai et al. (1998)

Porous coating without Liu et al. cracks, ethanol-based (2001) synthesis: thermally stable HA phase; water-based synthesis: calcium deficient HA phase; ~200 nm crystal size

11

12

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13

14

Calcium nitrate and triehtylphosphate in ethanol and distilled water Calcium nitrate and triethylphosphate in ethanol Ca-salt and triethylphosphate in ethanol and distilled water Calcium diethoxide and triethylphosphate in ethylene glycol, anhydrous ethanol and diethylene glycol

Commercial grade Ti

Drying at 80°C, final heating at 500°C

Ti6Al4V, poroussurfaced implants Ti

500°C

Ti

Conventional heat treatment combined with ultraviolet (UV) irradiation at ambient temperature

350°C

Rough, porous, pure HA ~5μm thick coatings, part of coating dissolved in 0.9% NaCl Uniform grain size and distribution; interfacial shear strength of 280 MPa Crack-free and porous coating

Lopatin et al. (1998)

HA crystallization at 500–7000°C; cytotoxicity decreased with increasing heat treatment temperature

Haddow et al. (1996)

Gan et al. (2004) and (2005) Liu et al. (2001)

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otherwise, calcium oxide also forms along with HA. Crack-free and homogeneous coatings are formed in most cases. Gan et al. (2005) and Gan and Pillar (2004), have used both inorganic and organic routes for sol-gel coatings of Ti6Al4V with HA. Ca/P molar ratio was maintained at 1.67 for both the methods. The solution pH was maintained at approximately 12 and 0.7 for the inorganic and organic methods, respectively. The substrates consisted of mill-annealed Ti6Al4V alloy which were either ground using 600-grit silicon carbide (SiC) paper or further polished using a 1 µm diamond suspension. Prior to coating, all the degreased and cleaned samples were finally given a standard nitric acid passivation treatment. The inorganic coating was five layers thick, whereas the organic coating was a single layer thick. A final annealing temperature of 500°C was used for the consolidation of the films followed by furnace cooling of the samples. The inorganic film was 1 µm thick and porous (60% density) with an irregular surface texture and had CaTi2O5 formed at the film/metal interface. The Ca/P ratio for the film was 1.46 and 2.10 for the inorganic and organic films, respectively. The high interfacial strength value, 347 MPa, was attributed to a primary chemical bond formed between the film and the substrate during the heat treatment. In a similar method of HA coating on Si (100) substrates by Hwang and Lim (1999), HA was formed at 500°C while β-TCP was formed at 700°C. In the study of Guo and Li (2004), the formation of nano HA lowered the temperature of crystallization to around 400°C. The crystal size increased and microstrain decreased with increasing final heat treatment temperature. The lower temperature of heat treatment resulted in non-stoichiometric HA. Cracks were more easily detected on the denser organic route-formed film (Gan et al., 2005), so that the strain for crack initiation is measured to be lower and leads to a lower value of the interfacial strength (280 MPa) – that is, a lower value measured does not necessarily imply an adhesion strength lower than that measured on the films prepared by the inorganic route. On the other hand, Liu et al. (2002) observed an apatite structure in the coatings annealed at temperatures higher than or equal to 400°C. Dense coatings with an average bonding strength of 44 MPa were formed. The difference in coating strength in these two investigations resulted from the different methods of testing. However, the bonding strength decreased in coatings annealed at 500°C due to the development of nanoporous structures as well as surface cracking in the coatings resulting from non-uniform thickness due to substrate roughness (sand-blasted). A contraction of the coatings resulting from sintering and phase transition may be responsible for the loss of structural integrity of the thicker sections of the coatings. Thus, dense and adhesive apatite coatings may be obtained through organic-route sol-gel technology after short-term annealing at 400–500°C in air. Kim et al. (2004) coated HA on to a titanium substrate with the insertion of a TiO2 buffer layer by the sol-gel method. A typical apatite phase

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was observed in the HA layer at 400°C and phase intensity increased above 450°C. The highest strength of the double-layer coating on titanium was 55 MPa after heat treatment at 500°C, which was ~60% greater than that of the HA single coating (35 MPa). The enhanced chemical affinity of TiO2 towards the HA layer as well as towards the titanium substrate and the dense and uniform coating structure resulted in improvement in bonding strength. Cell function and activity was enhanced at an early stage of cell differentiation.

5.6.3 Composite HA coatings by sol-gel deposition Several investigators have synthesized composite HA coatings and evaluated the properties. Weng et al. (2003) carried out in vitro evaluations of HA and fluorapatite/HA solid solution (FHA) coatings on Ti6Al4V obtained by a sol-gel route in Kokubo’s SBF and citric acid modified phosphate buffer solution (CPBS). The FHA film showed good bioactivity and better stability in CPBS and higher adhesion strength than the HA film. When F is incorporated into the film, crystallinity of the apatite film increases, and a decrease in intrinsic solubility of the FHA could significantly improve the stability. Moreover, the thermal expansion coefficient of the FHA film is closer to that of Ti6Al4V. Milella et al. (2001) prepared a TiO2/ HA composite coating on titanium substrates from a sol containing a mixture of TiO2 colloidal particles and HA submicron particles. Coating characterization showed that the low thickness coating was chemically clean, homogeneous, rough, and porous, with a welldefined phase composition as well as good adhesion strength of ~40 MPa. The composite coatings prepared by Im et al. (2007) were heat treated at 500°C for 2 h. The debonding strength of the coating (tested with a scratch tester) from a titanium substrate increased with increasing amount of TiO2, and was significantly higher when the amount of TiO2 exceeded 30%. Needle-shaped bone-like apatite formed after two weeks of exposure to SBF proved that the hybrid coatings were bioactive. The attachment density of cells also increased with the increasing amount of TiO2 and was significantly higher when the amount of TiO2 exceeded 50%. Roest et al. (2011) investigated the mechanical properties and adhesion behavior of sol-gel-derived HA nanocoatings on commercial purity titanium and Ti6Al4V. The HA-nanocoated Ti6Al4V substrates performed better than commercial purity titanium in terms of adhesive bonding. Specific anodizing treatments improved the adhesive bonding of HA to the substrates, most significantly in the titanium alloy. The toughness of the anodized layer is more than that of the HA layer for both the substrates. Hukovic et al. (2003) investigated the electrochemical corrosion properties of titanium, Ti6Al6Nb and Ti6Al4V coated with CaP coatings in HBSS.

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CaP coatings deposited by a sol-gel technique on titanium and a Ti6Al4V alloy were well crystallized mixtures of β-TCP and HA with some carbonate incorporated in HA. On the other hand, more amorphous HA and β-TCP phases were observed in the same coatings on a Ti6Al6Nb alloy. Here, HA contained more carbonate groups. Crystalline HA and β-TCP exhibited a beneficial corrosion protection effect on the substrate during prolonged exposure to HBSS. The authors (Choudhury and Agrawal, 2011) have formed pure HA and composite HA/ZrO2 coatings on commercial purity titanium and determined the interfacial shear strength and fretting wear resistance. The low values of the coefficient of friction (0.4–0.5) in SBF medium and morphology of the wear pits for considerably long cycles of fretting indicate strong bonding of the HA coating to the titanium surface. The maximum interfacial strength was found to be ~570 and 678 MPa, for the pure HA and composite films, respectively, on the polished and acid-etched surfaces. However, the maximum interfacial strength was found to be about 263 MPa on the oxidized surface. The steady-state crack spacing on these surfaces is shown in Fig 5.4 while that on the oxidized surface is shown in Fig 5.5. The relatively rough surface in the former (evident from the underlying grinding lines) plays a dominant role in strong adherence of the film with the substrate and increasing the interfacial shear strength of the HA coatings as compared with that of the highly polished surface. The presence of ZrO2 further enhances the coating/metal bond strength which is evident from the higher interfacial shear strength as well as the formation of localized cracks on the surface (Fig. 5.4b). The derivation of interfacial shear strength, τmax, is based on the assumption that the coating is free of residual stress. For sol-gel coatings, capillary stresses occurring during drying and annealing of the films results in residual tensile stresses (or strains), which later assist in crack development at lower applied strains during shear lag testing resulting in a reduction in the measured failure strength σf and failure strain ɛf. The higher CTE of HA (14 × 10−6/K) than that of Ti (8.6 µm/m/K) will result in residual thermal stresses contributing towards residual tensile strains and stresses in the film. The values obtained by Gan et al. are much lower when compared with our values obtained for the polished surface. The difference is due to the differences in substrate conditions. The value obtained for the oxidized surface in our case is comparable to that reported by Gan et al. (2005). Thus, it is evident that mechanical interlocking in the rougher surfaces (acid-etched) is responsible for the stronger bonding of the coatings with the substrate. Figures 5.6 and 5.7 show the variation in coefficient of friction (COF) of the coated surfaces and the morphology of some worn surfaces, respectively. Even though the COF is higher on passivated surfaces, the wear pit in this case appears less damaged (no cracks) than the oxidized surfaces. Thus, delamination is the primary mechanism of wear in this case.

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(a)

(b)

5.4 Steady-state crack spacing of (a) HA- and (b) HA + ZrO2-coated samples on passivated surfaces (Choudhury and Agrawal, 2011).

The surface roughness created by passivation treatment of titanium results in stronger bonding with the HA coating leading to better wear resistance. The values obtained for interfacial shear strength also prove the effectiveness of higher surface roughness in inducing stronger bonds between the titanium substrate and the HA coating. It can be inferred from all the previous sections, that strong adhesion between the substrate and coating plays an important role in the successful performance of the coating material. Table 5.5 summarizes the adhesion strength obtained by different researchers using various coating methods. It may be observed from the table that coatings of lower thickness in general have better adhesion with the substrate.

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5.5 Steady-state crack spacing on oxidized HA-coated sample (Choudhury and Agrawal, 2011).

0.55 0.50 0.45 Coefficient of friction (COF)

112

0.40 0.35 0.30 0.25

A

0.20

B

0.15

C

0.10

D

0.05 0.00 0

20000

40000

60000

80000

100000

Time (s) A: HAp + ZrO2 coating on oxidized surface B: HAp coating on oxidized surface C: HAp coating on passivated surace D: HAp + ZrO2 coating on passivated surface

5.6 Variation of COF with time for pure HA and composite coatings (Choudhury and Agrawal, 2011).

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(a)

(b)

5.7 Morphology of HA-coated worn surfaces (a) passivated and (b) oxidized surfaces, arrows indicating direction of oscillatory motion ((b) Choudhury and Agrawal, 2011).

5.7

Miscellaneous deposition techniques for HA coatings

5.7.1 Electrodeposition Coating by an electrodeposition method is carried out in an electrolyte cell where the substrate forms the anode. The electrolyte is composed of the salts of both calcium and phosphorus (Ban and Maruno, 1993, 1994; Shirkhanzadeh et al., 1994). In this method, it is possible to control the composition and coating structure by controlling the temperature during deposition or varying the current densities (Ban and Maruno, 1995). It is also possible to coat irregular surfaces.

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Table 5.5 Summary of adhesion strength obtained by different investigators using different coating methods

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No

Coating technique

1

Plasma spray

2

Plasma spray

HA+Ti6Al4V composite coating

∼28

Dong et al. (2003)

3 4

Biomimetic Sol-gel

9.5 347

Jonasova et al. (2002) Gan and Pillar (2004)

5

Sol-gel

280

6 7

Sol-gel Sol-gel

44 55 and 35

Gan and Pillar (2004) and Gan et al. (2005) Liu et al. (2001) Kim et al. (2004c)

8 9

Sol-gel Sol-gel

∼40 40

Milella et al. (2001) Kar et al. (2006)

10

Sol-gel

570, 263, and 678

Choudhury and Agrawal (2011)

11

Dip-coating from powder suspension Pulsed laser deposition RF sputter deposition

Alkali pretreated Ti Inorganic route, 1 μm thick Ground with 600 grit SiC paper or mirror polished Organic route, 1.5 μm thick Ground with 600 grit SiC paper or mirror polished Water-based route TiO2 buffer layer, double layer, and single layer TiO2 + HA composite Anodization and alkaline pretreatment HA (polished), HA (oxidized) and HA + ZrO2 (polished) 25 μm thick Ti6Al4V abraded with SiC papers

30

Mavis and Tas (2000)

30–40

Wang et al. (1997)

59.6

Zhao et al. (2006)

12 13

Coating variables/ thickness

Substrate conditions

Adhesion strength (MPa)

Gas-nitriding + acid etching and gas-nitriding+grit blasting

HA+(ZrO2+Y2O3), 3–4 μm

References Gross et al. (1998b)

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Manso et al. (2000) have used an electrodeposition method to study the effect of using different initial solutions, viz., calcium acetate and acetic acid, and sodium phosphate and sodium hydroxide, respectively. The coatings so obtained in aqueous basic conditions have been found to be polycrystalline HA and were strongly adherent to the substrate. Some of the Ca+ were substituted by Na+ in the structure resulting in a lower Ca/P ratio with respect to the stoichiometric sample. Like all other coating techniques, surface pretreatment also plays an important role in the electrodeposition coating method. Alkaline pretreatment helps in the nucleation of HA on the substrates with improved bond strengths. The Ca/P ratio and the morphology of the coating are sensitive to the alkaline surface treatment. Kar et al. (2006) reported an innovative method for preparing a nanotubular TiO2 surface by anodization at a low pH, and subsequent alkaline treatment before electrodeposition of HA nanocrystalline coating. The alkaline pretreatment provided a template for nucleation of the HA inside the nanotubes and increased the bond strength of the coating. The treatment in 0.5 M NaOH at 50°C for 2 min resulted in better surface morphology. The adhesion strength was further improved by annealing the HA-coated nanoporous TiO2 at 400–6000°C for 30 min in an argon atmosphere. Bond strength of 40 MPA was obtained after heat treatment at 600°C for 30 min in argon atmosphere. In a study by Abdel et al. (2008) in the alkaline pretreated surfaces, the coating thickness and weight gain decreased with increasing current density. On the untreated surfaces, a reverse trend was observed. The surface morphology changed from plate-like crystals on the untreated surfaces to needle-like crystals in the pretreated surfaces. The surface treatment also changed the Ca/P ratio from 1:1 to 1.667 – that is, closer to the stoichiometric value. The effect of a precoating treatment on the biomimetic and electrochemical deposition on porous titanium was compared by Zhang et al. (2003). Precoating treatments were done by acid etching (H2SO4 + HCl) and in some cases acid + alkali etching (in NaOH). Biomimetic deposition was done in a supersaturated solution containing Ca+2 and PO4−2, while the same solution was used as an electrolyte for electrodeposition. Biomimetic deposited coatings were found to be relatively uniform; however, electrodepositon was less sensitive to the pretreatment conditions of the titanium substrate, but produced less uniform and thinner coatings on the inner pore surfaces. The deposited layer consists of crystalline and flake-like octacalcium phosphate in both methods.

5.7.2 Electron beam deposition In the electron beam deposition technique, an electron beam is used to evaporate the material to be deposited. High-energy electrons generated in an

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electron gun are directed towards the surface of the source material kept in a crucible. The characteristics of the HA coating layer on the Ti surface formed by electron beam deposition have been reported by Kim et al. (2003). The Ca/P ratio of the coating layer had a strong influence on its dissolution behavior. Stability in the physiological medium was higher for coatings with a Ca/P ratio close to that of crystalline HA. Heat treatment of the coated layer in air at 400°C and 500°C improved the stability even further while increasing the interfacial bond strength with the substrate. Results of preliminary in vivo tests indicate that improvement in stability of the coatings due to heat treatment is beneficial for bone attachment to the implants.

5.7.3 Dip-coating from powder suspension Mavis and Tas (2000) have used chemically precipitated HA powders for coating Ti6Al4V substrates by a dip-coating method. Polyethylene glycol, glycerol and/or gelatin were the organic chemicals used for the powder synthesis. Calcium nitrate tetrahydrate and diammonium hydrogen phosphate were used as the sources of calcium and phosphorus. The substrates of Ti6Al4V were abraded with silicon carbide papers followed by washing with distilled water. A special apparatus was used for the dipping process which provided constant dipping and withdrawal speeds (e.g., 100 mm/min). 25 µm thick uniform HA coatings could be obtained by calcining the coated samples in a nitrogen atmosphere at 840°C. The thickness of the coatings could be varied by changing the solution, HA concentration and the rate of dipping/withdrawal. Highly porous HA coatings were obtained with bonding strengths of 30 MPa.

5.7.4 Polymeric route for HA coating A polymeric route for the formation of HA coatings has been suggested by Shibata et al. (2003) and Brendel et al. (1992). A mixed acetone solution of hydrolyzed phenyldichlorophosphine (C6H5PCl2) with Ca(NO3)2 is oxidized by bubbling with air. Polymerization reactions occur leading to an increase in viscosity. A Ti substrate is dipped into the solution when it becomes viscous. The film so formed is dried, followed by calcinations at 1100°C under vacuum, and the resultant HA coating has a thickness of about 2 µm, is porous, with a minor CaTiO3 phase.

5.7.5 High-velocity oxy-fuel (HVOF) spraying High-velocity oxy-fuel (HVOF) spraying is a dry process that produces a dense coating. HVOF thermal spray uses a fuel (i.e., propylene, hydrogen, propane, kerosene)/oxygen mixture in a combustion chamber. A simple

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Expansion nozzle

Powder in

Fuel gas

Oxygen

5.8 Schematic illustration of HVOF combustion gun.

schematic of the HVOF combustion gun is shown in Fig. 5.8. This combustion process melts HA particles in the powder that is continually fed into a gun using a carrier gas (argon) and propels it at high speeds (3000–4000 fps) towards the substrate. The high speed of the spray produces a dense coating upon impact. The effect of various processing conditions such as initial HA powder size, content of bioinert ceramic additives or annealing temperatures on HVOF spray technique of depositing HA has been evaluated by various investigators. Khor et al. (2003a, 2003b) used the HVOF spray technique to produce HA-based bioceramic coatings onto titanium alloy substrates. Processing conditions influence the structure of resultant coatings and, hence, their mechanical properties. The different melt state of the HA powders during coating deposition has a significant influence on the microstructure. Complete crystallization of the amorphous phase was observed by Li et al. (2002) at ~700°C and the crystallization temperature was dependent on the sample heating rate in the differential scanning calorimetry (DSC) test. The incorporation of ceramic additives like yttria stabilized zirconia (YSZ) and titania or annealing of the coatings at 750°C showed improvement in mechanical properties of HA coatings. However, chemical reactions between zirconia and HA or titania and HA could not be avoided during the coating formation. While the chemical reaction layer improves the structure and mechanical properties, the products detrimentally affect the biological performance of the coatings. The mechanism of the precipitation process in HVOF-sprayed HA coatings during exposure to SBF solution was studied by Khor et al. (2003b). The dissolution of TCP, TTCP or ACP phases at the coating surface increases the rate of precipitation of the dense, bone-like apatite due to a significant increase of Ca2+ in a localized area. The nanoindentation method determined the Young’s modulus of the precipitated layer to be ~120 GPa. The rate of precipitation was found to be directly dependent on the Ca2+

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concentration in the SBF near the coating surface and was found to be a partially diffusion-controlled process. In the study by Haman et al. (1995), no significant change was observed in the structure of coatings after immersion in a Ca2+ free HBSS.

5.7.6 Pulsed laser deposition In pulsed laser deposition, a highly intense laser beam is focused on a target where the high energy density during the laser pulse (about 1 GW within 25 ns) ablates almost any material (Bagratashvili et al., 1995; Tucker et al., 1996; Wang et al., 1997). The ablated material forms plasma, which is deposited on a substrate opposite the target. This method is quite flexible in preparing films under a wide range of deposition conditions such as kinetic energy, deposition rate and ambient gas (Bagratashvili et al., 1995; Jelinek et al., 1995). Wang et al. (1997) performed the structural and chemical characterization of the HA coating on Ti substrates obtained by pulsed laser deposition technique. The target was made of commercial HA. The use of argon-water vapor mixture with a pressure of 4.3 × 10−1 torr and substrate temperature in the range of 500–6000°C led to the formation of pure, uniform, crystalline HA films roughly 10 µm thick. The morphology of the coatings comprised of numerous spheroidal shaped particles of different sizes as well as columnar and dome-shaped structures. This process did not induce significant changes in the behavior of hydroxyl or phosphate functional groups. The adhesion strength of the coating was found to be ~30–40 MPa and linked to the fractography of the tested specimen. After the pulsed laser deposition, the Ca/P ratio in the HA film was found to have increased to 1.99. The Ca/P ratio was significantly higher, particularly near the coating/substrate interface, than in the coating interior.

5.7.7 Radio-frequency sputter deposition The radio-frequency (RF) deposition technique consists of coating under high vacuum conditions accompanied by the formation of plasma (Dijk et al., 1995, 1996; Wolke et al., 1994). In this process, a large area (anode) can be coated uniformly without the need for substrate motion. The plasma is initiated between the electrodes at pressures in the mTorr range by the application of a high voltage, either direct current or radio-frequency. The gas (argon in most cases) ions bombard the surface of the target producing a continuous stream of sputtered atoms that deposit on the nearby substrate. The plasma is sustained by the ionization caused by secondary electrons emitted from the cathode (HA target). The secondary electrons generated

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by ion bombardment are accelerated into the plasma. Argon is the most suitable process gas for a good deposition rate and low cost. Plasma-sprayed HA was used as the target material by several investigators (Djik et al., 1995). Crystallite size and crystallinity increased with increasing temperature of heat treatment. Annealing at more than 600°C resulted in crystalline HA (Djik et al., 1996). The high Ca/P ratio in the as-sputtered coatings decreased with increasing temperature. The rate of deposition and Ca/P ratio increased with increasing argon pressure. With increasing annealing temperature, the second phase in the coating changed from TTCP to calcium oxide to β-TCP. Crystallite size increased significantly after immersion for one week in α-microelectro mechanical (MEM), a balanced salt solution usually used for cell culture studies. The increase in crystallinity during immersion was possibly due to dissolution of the amorphous phases. The dissolution of the coatings may have resulted in supersaturation of Ca and P ions in the physiological media causing reprecipitation of a crystallized phase. Carbon concentration in the coatings increased after one week of immersion in solution which is due to the incorporation of carbonate from the solution into the coatings. Zhao et al. (2006) have successfully developed 3–4 µm thick HA + (ZrO2 +Y2O3) composite coatings on Ti6Al4V substrates, using RF magnetron sputtering technique. The surface morphology of the coating was rough and uneven with 30–40% of the surface being covered with concave and convex pits, as shown in Fig. 5.9. The presence of these pits would be helpful for the growth of tissues when used as biological implants. The Ca/P ratios in the assputtered and post-annealed coatings were both in the range of 1.77~1.79. The ultimate bond strength of the sputtered composite coatings, 59.6 MPa,

5.9 SEM surface morphology of HA + (ZrO2 + Y2O3)/Ti6Al4V composite coatings fabricated by RF magnetron sputtering (Zhao et al., 2006).

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increased with the increasing sputtering power and the addition of YSZ. However, the pure HA coatings exhibited a maximum strength of ~55 MPa. No cracks/pores were observed in the coating/substrate interface. The high bond strength of the composite coatings was attributed to the formation of a diffusion layer (1.5–2 µm thick) in the interface, and the mechanical interlocking resulting from the rougher surface of the substrate treated by sand blasting and etching.

5.8

Conclusions

The poor mechanical properties of bulk HA led to the development of HA as coating on metallic implants. By using HA-coated implants, advantage is taken of the mechanical benefits of the underlying metals as well as the biocompatibility of HA. HA can bond directly to the bone, which can achieve higher fixation strength; it also prevents metallic corrosion. Since the mid1980s, plasma spraying has been used for the coating of HA on metallic implants and has now developed into the most widely used coating method. The more recently developed and promising methods of coating HA are sol-gel and biomimetic processes. Both these methods have the advantages of low processing temperatures, ability to coat complex shapes and elimination of the need for elaborate equipment. Extensive research has been conducted on the effect of substrate condition, precursors/feedstock material and processing variables on the properties of the coatings. The most widely studied properties of the coatings are the microstructure, chemical homogeneity and crystallinity, adhesion strength and dissolution behavior in different liquid media. The performance of the HA coating largely depends on its structural integrity which in turn depends on the processing conditions. Much of the present review concentrates on the effect of various processing parameters on the quality of the coating produced by different methods particularly, plasma spraying, sol-gel and biomimetic processes. The effect of plasma-spraying variables, powder feedstock size and morphology on the microstructure and mechanical properties of the plasmasprayed coatings has been extensively studied. The most widely studied mechanical properties of plasma-sprayed coatings are microhardness, Young’s modulus, fracture toughness, bond strength and wear. Tensile adhesion strength is strongly affected by the substrate pretreatment and the presence of reinforcements in the coatings. Bond strength of the coating/implant interface, particularly in a physiological environment, is an important criterion for the in vivo stability of the coatings and should be a focus for future research work. The chemical stability of the coatings in a physiological environment is important for the long-term use of the implants.

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The choice of precursors and solvents influences the nature of the HA coating prepared by sol-gel methods. The heat treatment temperature significantly affects the crystallinity and morphology of the coating, which in turn determines the behavior of the coatings under different conditions. Higher temperature generally results in more crystalline HA but at the same time increases the chances of degradation of the substrate material. Different surface treatments, such as nitric acid passivation and grit blasting, have been used to improve the adherence of sol-gel coatings. Further work to assess the stability and bioactivity of HA coatings prepared by sol-gel process is recommended. Alkali pretreatment is the most commonly used method to modify the surface characteristics suitable for biomimetic coating of sol-gel. Acid treatment combined with alkali pretreatment has also been found to be useful for the biomimetic deposition of HA. It is necessary to assess the effect of various other surface treatments on the various properties of biomimetic HA coatings. A number of other deposition methods of HA coatings have also been reviewed here. Most of these methods are still in the initial stages of development as far as HA coatings are concerned.

5.9

Future trends

Implant materials of today must be designed such that they last as long as the lifetime of the patient. An ideal implant must have a suitable surface chemistry that causes cell changes to occur at the interface which would normally occur in the absence of the implant. Non-adherent fibrous capsules are formed on implants which do not satisfy such conditions, resulting in their interface instability. Infection, continuous wear and release of worn particles, release of toxic substances or uncontrolled surface degradation leads to formation of a fibrous capsule that isolates the implant from the normal tissues. In other words, an ideal implant is one which behaves in a similar way to host tissue. In the light of the above, the focus of future research must be on creating an implant-tissue interface which is simultaneously histologically and biomechanically stable. The short-term response of the tissues in the presence of implants determines the eventual long-term response. Therefore, for all future research on HA, emphasis must be laid on cell culture experiments. Such experiments help to determine the behavior of different types of cells on various implant surfaces. Highly chemically specific attachment mechanisms control cell morphology, influence the structure of cell membranes and activate processes required for differentiation. Bioactive implant surfaces are generally helpful in the growth and attachment of certain cells (primary osteoblast-like cells), whereas, such surfaces

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slow down the attachment, spreading and growth of fibroblast cell lines. Post-formation surface treatment is necessary for the HA coatings so as to modify the surface activity according to the cell requirement.

5.10

Acknowledgement

Financial assistance received from the Department of Science and Technology (DST), Government of India, is gratefully acknowledged.

5.11

References

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