Wear properties of micro arc oxidized and hydrothermally treated Ti6Al4V alloy in simulated body fluid

Wear properties of micro arc oxidized and hydrothermally treated Ti6Al4V alloy in simulated body fluid

Wear 271 (2011) 2322–2327 Contents lists available at ScienceDirect Wear journal homepage: www.elsevier.com/locate/wear Wear properties of micro ar...

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Wear 271 (2011) 2322–2327

Contents lists available at ScienceDirect

Wear journal homepage: www.elsevier.com/locate/wear

Wear properties of micro arc oxidized and hydrothermally treated Ti6Al4V alloy in simulated body fluid Y. Vangolu a , A. Alsaran a,∗ , O.S. Yildirim b a b

Department of Mechanical Engineering, Faculty of Engineering, Atatürk University, 25240 Erzurum, Turkey Department of Orthopaedics, Faculty of Medicine, Atatürk University, 25240, Erzurum, Turkey

a r t i c l e

i n f o

Article history: Received 2 September 2010 Received in revised form 1 December 2010 Accepted 1 December 2010

Keywords: Bioceramics Hydroxyapatite Micro arc Hydrothermal treatment Wear SBF

a b s t r a c t Hydroxyapatite (HA) has a strong ability to form a direct chemical bond with the hard tissues. It is also synthesized on Ti6Al4V alloy using micro arc oxidation (MAO) followed by hydrothermal treatment (HT). In this study, Ca and P containing titania coated surfaces were achieved from the MAO coating process which was carried out at 450 V for 5 minutes and at a frequency of 300 Hz in a calcium acetate monohydrate and ␤-glycerophosphate aqueous solution. The substrates coated with TiO2 were hydrothermally treated inside a NaOH solution which had the pH value of 11.0–11.5 at 160 ◦ C, 190 ◦ C and 220 ◦ C for 5 h and 10 h in a pressure vessel. Microstructure and phase components of the coatings were characterized using scanning electron microscope (SEM) and X-Ray diffraction (XRD) before and after micro arc oxidation and hydrothermal treatment. Each untreated, Ca and P containing TiO2 coated and HA coated specimens were subjected to pin on disk wear tests inside simulated body fluid (SBF) at 37 ◦ C. It was observed that after HT, rough and porous bioactive HA crystals were formed on the surface. HA/TiO2 film layer formed with MAO plus HT improved the wear properties of Ti6Al4V. By this way, a surface which has a good wear resistance and bone bonding ability was achieved. © 2011 Elsevier B.V. All rights reserved.

1. Introduction Titanium alloys are potentially very suitable materials for load bearing bio-implant applications because of their good ductility, formability, high modulus of elasticity and strength to weight ratio. However the applications of titanium and its alloys are limited because of their low surface hardness, high friction coefficient, poor wear resistance. Besides, their bioactivity and biocompatibility with osseous tissues are not as good as certain forms of calcium phosphate ceramics such as hydroxyapatite (HA, Ca10 (PO4 )6 (OH)2 ). Because of easily oxidizable nature of Ti, several oxidation processes have been applied to Ti and Ti alloys in order to improve their wear and corrosion resistance. Besides, this oxide layer, which is easily formed under moist conditions, acts as a chemical protective barrier to obstruct the metal ion liberation from the implant in in vivo media (pH level of 7.4 and a temperature of 37 ◦ C). The TiO2 layer should be thick enough to provide effective corrosion resistance in body fluid. Consequently this type of materials must sustain their wear resistance and biocompatibility in biological environment. Recently there are several surface treatments applied to improve these requirements and HA/TiO2 composite coatings

∗ Corresponding author. E-mail address: [email protected] (A. Alsaran). 0043-1648/$ – see front matter © 2011 Elsevier B.V. All rights reserved. doi:10.1016/j.wear.2010.12.039

appear to be one of the most suitable way to achieve both solicited properties together [1–5]. Hydroxyapatite (HA) is an important bio-ceramic material which has a high biocompatibility and as a result HA can improve the bone bonding ability of the implant surface. However, practical clinical applications of HA under load-bearing conditions have been limited due to the poor mechanical properties of bulk HA ceramics [6]. So, an effective method to solve the above mentioned problems is to use the hydroxyapatite as a coating on the surface of Ti and Ti alloys. Many methods can be used to deposit HA especially on Ti surfaces such as plasma spraying, electrophoretic deposition and dip coating [7–9]. Currently, plasma spraying is the most popular technique for HA coatings. Many studies have shown that plasma sprayed HA coatings improve the biocompatibility, integration and also promote the bone attachment of the implant [10–13]. Because of its clinical success plasma spraying became the only commercially feasible way to deposit HA on the implant surfaces. However, the uncontrollable dissolution tendency, unsuitable coating nature (for complex geometries), poor fracture toughness, maintaining phase purity, mechanical strength in vivo and bond strength of plasma sprayed coatings are among the problems for long-term stability [6,14–16]. Many researches have been carried out to improve these properties of HA coatings. In the last five years, micro arc oxidation followed by hydrothermal treatment were introduced for HA/TiO2 coatings. MAO is one of the effective methods to form

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Table 1 The mechanical properties of hydroxyapatite. Mechanical Properties Tensile strength, ultimate Modulus of elasticity Flexural strength Compressive strength Poissons ratio

38.0–48.0 MPa 7.00–13.0 GPa 100–120 MPa 350–450 MPa 0.270

porous, rough and homogeneous oxide layers, which can be controlled by process parameters, on the surface. Furthermore, Ca and P elements which are the main components of hydroxyapatite can be merged into the titania layer by changing the electrolyte. The formation phenomena of hydroxyapatite crystals via micro arc oxidation and hydrothermal treatment were explained by Li et al. [17], Ishizawa et al. [18–20] and Liu et al. [21]. With this method, better adhesion of HA to the substrate was achieved along with the fact that improvement of the mechanical properties, corrosion and wear resistance and biocompatibility were obtained by forming a HA/TiO2 composite layer on the surface [1,14,22–27]. However, few investigations on wear resistance behaviors of HA/TiO2 coating have been carried out in the scientific literature. The main purpose of this work is to understand the effects of micro arc oxidation and hydrothermal treatment on morphology and in vitro wear properties of Ti6Al4V alloy. The effects of hydrothermal treatment temperature and time have also been investigated. Hydroxyapatite layers were formed on the surface of Ti6Al4V substrates by MAO and HT. It is believed that this method can improve the adhesion and the wear rate of the HA coatings which is a significant problem of plasma spraying technique and it can be an alternative method to plasma sprayed HA coatings in commercial area. The results obtained from the tests are promising and it is worth further investigation for the use in orthopedic and dental applications. Structural and bio-tribological properties of the coatings were characterized by using XRD, SEM/EPMA and pin-on-disc tribotester. 2. Experimental Ti6Al4V (Grade 5) plates with the dimensions of 20 mm × 20 mm × 0.3 mm were used as the substrates in MAO process. The substrates were mirror-like polished and then rinsed by HF+HNO3 +HCl+Ethanol mixture. The specimens were cleaned with acetone, alcohol and distilled water. The MAO process was carried out in an electrolyte which consists of 0.2 M Ca(C2 H3 O2 )2 ·H2 O (Calcium acetate monohydrate), 0.02 M C3 H7 CaO6 P.aq (Calcium ␤-glycerophosphate) and distilled water. For the treatment, unipolar pulse voltage was applied by an AC pulse power supply while the specimens were the anode of the system and a stainless steel bath was the cathode. The system temperature was kept constant during the experiments via a circulation pump. The MAO was performed at a constant voltage, frequency, duty circle and time of +450 V, 300 Hz, 6–6.5% and 5 min. respectively. The coated samples were flushed with water after MAO treatment and then they were placed inside a 100 ml aqueous solution in an autoclave with the pH value of 11.0–11.5 for time intervals 5 h and 10 h and temperatures of 160 ◦ C, 190 ◦ C and 220 ◦ C in order to form hydroxyapatite. The pH was adjusted via NaOH addition. The mechanical properties of HA were given in Table 1. Conventional scanning electron microscopy (SEM/EDS) was employed to observe the morphology on the surface. Film phase structures were obtained using a X-ray diffractometer (XRD) by scanning at 30 kV and 30 mA with CuK␣ radiation. Wear experiments were performed in SBF solution at 37 ◦ C using a pin-on-disc tribotester. All wear experiments were conducted

Fig. 1. The XRD spectra of the specimens coated by MAO and hydrothermally treated for 5 h.

with 6 mm diameter tungsten carbide (WC) balls in sliding contact with a load of 1 N at a sliding velocity of 6.5 cm/s. Wear tests wear carried out for a sliding distance of 50 m. Coating adhesion was investigated using scratch tester in which a diamond stylus is drawn across a surface with increasing normal load until well defined failures occur, at a certain critical load (Lc). The critical load (Lc) values were measured with a 0.2 mm tip radius Rockwell C diamond indenter, under loading rate of 100 N min−1 and sliding speed of 10 mm min−1 . The critical load values are detected with optical microscope by examination of the scratch track. 3. Results and discussion 3.1. XRD analyses The XRD patterns of Ti6Al4V after MAO and HT are given in Figs. 1 and 2. Ti6Al4V has a phase structure of ␣-Ti and ␤-Ti. After MAO and HT, a composite structure consisted of rutile, anatase and hydroxyapatite (HA) phases was formed on the surface of Ti6Al4V. Liu et al. [21] and Khan et al. [28] acquired similar results in their studies. The rutile and the anatase phases, which are formed after MAO process, are the two major polymorphs of TiO2 . Li et al. [17] indicated that the porous anatase induces HA nucleation. Although the MAO process carried out in an electrolyte containing Ca and P elements, no intensities connected to Ca and P elements were detected from the XRD patterns. Nevertheless it is accepted that the Ca and P elements exist in amorphous TiO2 layer [20,21,28,29]. After HT, HA peaks were distinctly determined on the surface layer. When the XRD spectra of the specimens in Fig. 1 examined, it was observed that as the HT temperature increased, the HA peaks became higher and the highest peak was observed at 220 ◦ C for

Fig. 2. The XRD spectra of the specimens coated by MAO and hydrothermally treated for 10 h.


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Fig. 3. The SEM micrographs of (a) Ca-P containing TiO2 layer formed by MAO (b) 160 ◦ C – 5 h hydrothermally treated and (c) 190 ◦ C – 10 h hydrothermally treated.

5 h treated samples, but the difference between the samples was quite small. While the intensities of the HA at 33◦ were significantly changed, the HA peaks between 45◦ and 55◦ were marginal as the temperature increased. The XRD spectra of the specimens treated for 10 h at various treatment temperatures versus CaP containing MAO is given in Fig. 2. HA peaks were clearly seen at both HT temperatures. When only the effect of HT temperature for 10 h treatment time was investigated, it was observed that the HA peaks of the sample treated at 190 ◦ C were higher than that of either 160 ◦ C or 220 ◦ C treated specimens. So by contrast with the samples treated for 5 h, HA peaks decreased after 190 ◦ C. It can be seen from Fig. 2 that the specimens treated at 220 ◦ C for 10 h, the intensity of HA peaks are lower. This indicates that the columnar structure of HA crystals may be altered and hence the surface morphology may be changed after this temperature with being held for a time. Besides, a discoloration of the oxidized surface was noted on the surface of the specimen treated at 220 ◦ C. Almost all expected HA peaks were seen at the XRD pattern of the sample treated at 190 ◦ C for 10 h. Even though the observed HA peaks were not quite different either for 5 h or 10 h treatment times at both HT temperatures, relatively the best XRD patterns were obtained from 190 ◦ C for 10 h treated sample. In addition to all these, Liu et al. [21] indicated that the most effective parameter on the nucleation and formation of HA crystals is the pH value of the water solution which is used for HT.

3.2. SEM/EDS analyses Fig. 3 shows the surface topography and EDS analyses of the HA/TiO2 formed on the surface of Ti6Al4V. A homogenous film layer which has a thickness of 10 ␮m was achieved after MAO and HT. As seen in Fig. 3(a), the surface morphology of the oxide layer after MAO treatment was rough and porous. The pores with miscellaneous dimensions including micro pores and sub-micro pores, which are a natural result of MAO, are related to the frequency value of the process. Lv et al. [30] reported that larger pore sizes and rougher coatings can be achieved at relatively low frequencies. In this study, despite the results of XRD analyses, it is known from the EDS analyses that the titania layer contains Ca and P elements, so the crystal-looking particles which can be seen in Fig. 3(a), are thought to be the Ca and P based compounds. No signs of HA discovered from the SEM images of the oxidized substrates. HT carried out in an autoclave filled with pH adjusted water solution that induces the HA nucleation [7,17,20,21]. So after hydrothermal treatment at various times and temperatures HA crystals were observed on the surfaces of the substrates. Small amounts of HA fibers were observed from the substrates hydrothermally treated at 160 ◦ C for 5 h (Fig. 3(b)). It can be said that the formation of HA was inadequate or the HA cannot fully grow on the surface. Relatively higher amounts of HA dendrites were discovered in patches at the specimens treated for 5 h at 190 ◦ C. But these dendrites

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Fig. 4. COF of various coatings and Ti6Al4V versus WC ball. Fig. 5. The Lc values and the wear rates of the samples.

did not cover the whole surface and no uniformity was noticed. When the treatment time extended from 5 h to 10 h, an increase of the extensity of HA was noticed. Liu et al. [21] indicated that the HT makes Ca and P diffuse from inner layer to the surface to compensate the deficiency in order to resume the reaction of HA formation. Thus, increasing the treatment time may enhance the diffused amount of Ca and P and so may the amount of HA. Actually HA amount acquired from all hydrothermal temperatures carried out for 10 h were higher than the ones treated for 5 h. It was observed that a large number of columnar hydroxyapatite crystals are precipitated on the surface of specimens hydrothermally treated at 190 ◦ C for 10 h (Fig. 3(c)). HA crystals completely covered the surface. Even the columnar HA crystals grew in through the pores. In this way the bone ingrowth to the implant, which is referred to be one of the important features that is expected from a biomaterial, becomes easier [31,32]. Ca/P ratio is also one of the important parameters in HA coatings. It is expected to have a minimum value of 1.67. In this study, Ca/P ratio of the micro arc oxidised layer, which contains Ca and P ions, was ranged between 1.70 and 3.64. 3.3. Wear and Scratch tests The graphics showing the variation of the coefficient of friction (COF) versus time for untreated, MAO processed and hydrothermally treated Ti6Al4V alloy are given in Fig. 4. At the beginning of the tests, the friction coefficient showed an increasing trend because of the Hertzian contact. After a certain time surface roughness decreased and as a result coefficient of friction became stable until the end of the tests. At the later stages of the tests, the friction coefficient increased or decreased according to the adhesion of the coated layer. The adhesion of the coated layer was relatively low on the specimens which the hydroxyapatite layer did not form properly. It can be seen in the figure that the friction coefficients were decreased in every specimen after the surface treatments. The relatively lowest friction coefficient (0.4) was obtained from the specimen hydrothermally treated at 220 ◦ C for 5 h, while the friction coefficient was approximately 0.6 for the specimen treated only with MAO. The friction coefficient was rapidly increased after 150 s on the specimens hydrothermally treated at 160 ◦ C and 190 ◦ C for 10 h. Fu et al. [33] determined that the HA coatings on Ti6Al4V formed with plasma spraying has a friction value of 0.75 in dry friction conditions, where the coefficient is 0.35 in bovine albumin media. Also, in their study, the friction coefficient remains stable until the test ends. For this reason the bio-simulated media may perform as a lubricant. In our study, an increase of the friction coefficient was observed after a certain time on the tests carried out in SBF solution. This result showed that the wetting capability of the SBF was low and the lubricant effect of the solution decreased with the increase of the wear time.

The relationship between the wear rate and the Lc values is given in Fig. 5. The wear resistance of all specimens increased compared to that of Ti6Al4V. It is known that the wear mechanism of Ti6Al4V starts with oxidation and then adhesive and abrasive wear takes place [34]. Titanium has a native oxide layer on its surface because of its high affinity to the oxygen. If this passive layer is stable, the wear resistance decreases. If it is not stable and the applied normal and shear stresses are high enough to disrupt this layer, then the abrasive effect takes place and the wear resistance decreases [35]. As seen in Fig. 6(a), deep wear tracks formed due to the abrasion of Ti6Al4V, and abrasive particles were observed on the inner surface of the wear tracks. The width of the wear track of TiO2 layer, which was formed by MAO, was narrow compared with the untreated Ti6Al4V. The wear mechanism for oxidized materials depends on the adhesion between the film layer and the substrate. So, the weak adhesion causes the disruption of the oxide layer. This layer may bond adhesively to the mating body or it also may modify a passive layer. The modified passive layer may be subjected to a continuous removal in the first situation. This continuous removal gradually depletes both the coating and the substrate and increases the roughness, causing a relatively high metal loss. Ultimately, the breakdown of the oxide layer leads to the potential for abrasive wear (Fig. 6(b)), where the hard oxide debris acts as third body abrasive components. The wear track width of HA/TiO2 composite layers is almost close to that of the specimens coated only with Titania layer (Fig. 6(c and d)). If the HA wear products transfer to the counter surface during the wear tests, a soft–soft surface contact may occur. Therefore both the friction coefficient and the abrasive effect can decrease. EDS analyses conducted after the wear tests did not reveal any trace of particles from the pin which might be transferred to the substrate surface. After wear, oxygen and titanium had the highest values in weight and only the amounts of Ca and P were reduced (Fig. 6(e and f)). The presence of the Ca and P elements on the surface of the specimen even after the wear tests showed that the coating technique increased the wear resistance of hydroxyapatite layer. It was noticed from the observations of the adhesion behaviour of HA coatings that the Lc values are varied between 5-10N in both coated specimens. The Lc is the value where the first damage (disruption) begins. The achieved values in this study were consistent with the results of Nie et al. [1]. The adhesion of the HA/TiO2 coatings are related to the structure of the coating. Generally a HA layer, which has a good crystal structure and high density, has good adhesion [40]. For this reason, the HT is an effective parameter on adhesion. From the XRD patterns, the relatively highest HA density was acquired from the specimens hydrothermally treated at 190 ◦ C. The observation of Lc values showed that the best adhesion conditions were obtained at the same temperature. Besides, the HT time effected the formation and the density of HA, so the adhesion increased with holding for a time.


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Fig. 6. The SEM micrographs of the wear tracks and the graphics of EDS analyses.

The adhesion of the HA layer directly effects the wear resistance. It is known that the most important disadvantage of the HA coating is its low resistance in long-term wear conditions and it is necessary to overcome this problem. The coating disrupts after a certain period of time because of the weak adhesion and the disrupted particles create abrasive effects. Nowadays there are several methods which are used to deposit HA but the main problem of these methods is the weak adhesion which is affected by the coating parameters. This problem prevents the use of the HA coatings as a wear resistant coating. For example, it is known that in plasma sprayed HA coatings, the dimensions of HA particles and its low bonding strength to the substrate decreases the wear resistance [6,10,36–39]. Actually the crystal form of the HA coating is the most effective parameter on the adhesion. MAO and hydrothermal treatment allow to form homogenous HA crystals on the surface which improve the adhesion. Thus, the increased adhesion makes the separation of the coating from the substrate difficult and the wear resistance increase. Using the scratch test, relatively high adhesive strength compared to other techniques as thermal spraying was found between the coating layers and the substrate [41].

4. Conclusions The following conclusions can be drawn from this study: • While the TiO2 layer formed on the surface of Ti6Al4V after MAO, HA containing titania layer was formed after MAO and HT. For the specimens hydrothermally treated for 10 h, the HA density increased (approx. 21%) with the increasing HT temperature. Although, the HA density began to decrease (approx. 29%) at HT temperatures higher than 190 ◦ C for 10 h treatment time. And for the 5 h treated specimens, the HA density increased approximately 38% as the HT temperature increased from 160 ◦ C to 220 ◦ C. It was observed that the specimens treated at 190 ◦ C for 10 h has the highest HA density. • Since there were no HA diffractions observed on the surface of micro arc oxidised samples from the XRD patterns, the EDS analyses indicated that the Ca and P elements were present in the structure and the Ca/P ratio of this layer was varied between 1.70 and 2.19. The hydroxyapatite peaks, which were not observed on the XRD patterns of MAO, occurred after hydrothermal treatment, indicating that the hydrothermal treatment induced hydrox-

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yapatite nucleation and formation on the surface of Ca and P containing titania layer. At the HT temperature of 190 ◦ C, HA was homogenously dispersed on the surface while it was not at lower HT temperatures. Besides HA crystals clustered at temperatures higher than 190 ◦ C. • It was found on the tests carried out in SBF that the wear resistance of all the coated specimens increased when compared with the untreated Ti6Al4V. The wear rate of the micro arc oxidised specimens improved about 75% and wear rate of the specimens which were hydrothermally treated at 190 ◦ C for 10 h after MAO improved about 82% compared with the untreated Ti6Al4V. TiO2 layer on the surface was disrupted as a result of wear which caused abrasive effect. The abrasive effect was relatively low on the surfaces which contain HA. According to the EDS analyses, the amount of Ca and P was decreased as a result of wear. • It was observed that the adhesion behaviour of the coating was directly related to the HT temperature. It is known from the scientific literature that a good crystallization of hydroxyapatite phase provides a good adhesion to the surface. The results showed that the crystallinity and the Ca/P ratio became better as the HT temperature reached 190 ◦ C. So, the best adhesion was obtained at 190 ◦ C – 10 h where the relatively dense HA crystals were formed. Besides the enhancement of the adhesion directly provided the increment of the wear resistance. • The adhesion and wear properties of HA coatings formed by MAO and subsequent HT were improved. This technique can be used as an alternative way to commercially deposit thin HA layers (max. 10–15 ␮m). Further investigations, especially in vivo are worth conducting for use in medicine. Acknowledgement The authors would like to thank Res. Asst. Fatih Erdemir from Materials and Metallurgy Faculty of Karadeniz Technical University for SEM and EDS analyses. Also, the authors would like to thank Prof. Dr. Ihsan Efeoglu for scratch tests. References [1] X. Nie, A. Leyland, A. Matthews, Deposition of layered bioceramic hydroxyapatite/TiO2 coatings on titanium alloys using a hybrid technique of micro-arc oxidation and electrophoresis, Surf. Coat. Technol. 125 (2000) 407–414. [2] J.M. Choi, H.E. Kim, I.S. Lee, Ion-beam-assisted deposition (IBAD) of hydroxyapatite coating layer on Ti-based metal substrate, Biomaterials 21 (2000) 469–473. [3] O.S. Yildirim, B. Aksakal, H. Celik, Y. Vangolu, A. Okur, An investigation of the effects of hydroxyapatite coatings on the fixation strength of cortical screws, Med. Eng. Phys. 27 (2005) 221–228. [4] A. Stoch, A. Brozek, G. Kmita, J. Stoch, W. Jastrzebski, A. Rakowska, Electrophoretic coating of hydroxyapatite on titanium implants, J. Mol. Struct. 596 (2001) 191–200. [5] A. Costantini, G. Luciani, F. Branda, L. Ambrosio, G. Mattogno, L. Pandolfi, Hydroxyapatite coating of titanium by biomimetic method, J. Mater. Sci.: Mater. M 13 (2002) 891–894. [6] F.J. GarciaSanz, M.B. Mayor, J.L. Arias, J. Pou, B. Leon, M. PerezAmor, Hydroxyapatite coatings: a comparative study between plasma-spray and pulsed laser deposition techniques, J. Mater. Sci.:Mater. M 8 (1997) 861–865. [7] B.C. Yang, M. Uchida, H.M. Kim, X.D. Zhang, T. Kokubo, Preparation of bioactive titanium metal via anodic oxidation treatment, Biomaterials 25 (2004) 1003–1010. [8] Y.M. Lim, Y.J. Park, Y.H. Yun, K.S. Hwang, Functionally graded Ti/HAP coatings on Ti–6Al–4V obtained by chemical solution deposition, Ceram. Int. 28 (2002) 37–41. [9] F. Liu, F.P. Wang, T. Shimizu, K. Igarashi, L.C. Zhao, Hydroxyapatite formation on oxide films containing Ca and P by hydrothermal treatment, Ceram. Int. 32 (2006) 527–531. [10] K. De Groot, R. Geesink, C.P.A.T. Klein, P. Serekian, Plasma sprayed coatings of hydroxylapatite, J. Biomed. Mater. Res. 21 (1987) 1375–1381. [11] L.M. Sun, C.C. Berndt, K.A. Gross, A. Kucuk, Material fundamentals and clinical performance of plasma-sprayed hydroxyapatite coatings: A review, J. Biomed. Mater. Res. 58 (2001) 570–592.


[12] R.Z. LeGeros, J.P. LeGeros, Y. Kim, R. Kijkowska, R. Zheng, C. Bautista, J.L. Wong, Calcium phosphates in plasma-sprayed HA coatings, Ceram. Trans. 48 (1995) 173–189. [13] X.B. Zheng, C.X. Ding, Characterization of plasma-sprayed hydroxyapatite/TiO2 composite coatings, J. Therm. Spray Technol. 9 (2000) 520–525. [14] Y. Chen, T.H. Zhang, C.H. Gan, G. Yu, Wear studies of hydroxyapatite composite coating reinforced by carbon nanotubes, Carbon 45 (2007) 998–1004. [15] J.F. Kay, Calcium phosphate coatings for dental implants. Current status and future potential, Dent. Clin. North. Am. 36 (1992) 1–18. [16] H.T. Zeng, W.R. Lacefield, XPS, EDX and FTIR analysis of pulsed laser deposited calcium phosphate bioceramic coatings: the effects of various process parameters, Biomaterials 21 (2000) 23–30. [17] P.J. Li, C. Ohtsuki, T. Kokubo, K. Nakanishi, N. Soga, K. Degroot, The role of hydrated silica, titania, and alumina in inducing apatite on implants, J. Biomed. Mater. Res. 28 (1994) 7–15. [18] H Ishizawa, M. Ogino, Characterization of thin hydroxyapatite layers formed on anodic titanium-oxide films containing Ca and P by hydrothermal treatment, J. Biomed. Mater. Res. 29 (1995) 1071–1079. [19] H. Ishizawa, M. Ogino, Thin hydroxyapatite layers formed on porous titanium using electrochemical and hydrothermal reaction, J. Mater. Sci. 31 (1996) 6279–6284. [20] H. Ishizawa, M. Ogino, Hydrothermal precipitation of hydroxyapatite on anodic titanium oxide films containing Ca and P, J. Mater. Sci. 34 (1999) 5893–5898. [21] F. Liu, Y. Song, F.P. Wang, T. Shimizu, K. Igarashi, L.C. Zhao, Formation characterization of hydroxyapatite on titanium by microarc oxidation and hydrothermal treatment, J. Biosci. Bioeng. 100 (2005) 100–104. [22] J. Baszkiewicz, D. Krupa, J. Mizera, J.W. Sobczak, A. Bilinski, Corrosion resistance of the surface layers formed on titanium by plasma electrolytic oxidation and hydrothermal treatment, Vacuum 78 (2005) 143–147. [23] Y. Han, S.H. Hong, K.W. Xu, Structure and in vitro bioactivity of titania-based films by micro-arc oxidation, Surf. Coat. Technol. 168 (2003) 249–258. [24] Y. Han, S.H. Hong, K.W. Xu, Synthesis of nanocrystalline titania films by microarc oxidation, Mater. Lett. 56 (2002) 744–747. [25] D.Y. Kim, M. Kim, H.E. Kim, Y.H. Koh, H.W. Kim, J.H. Jang, Formation of hydroxyapatite within porous TiO2 layer by micro-arc oxidation coupled with electrophoretic deposition, Acta Biomater. 5 (2009) 2196–2205. [26] D. Lahiri, V. Singh, A.K. Keshri, S. Seal, A. Agarwal, Carbon nanotube toughened hydroxyapatite by spark plasma sintering: microstructural evolution and multiscale tribological properties, Carbon 48 (2010) 3103–3120. [27] J.E. Tercero, S. Namin, D. Lahiri, K. Balani, N. Tsoukias, A. Agarwal, Effect of carbon nanotube and aluminum oxide addition on plasma-sprayed hydroxyapatite coating’s mechanical properties and biocompatibility, Mater. Sci. Eng. C: Mater. 29 (2009) 2195–2202. [28] R.H.U. Khan, A.L. Yerokhin, A. Matthews, Structural characteristics and residual stresses in oxide films produced on Ti by pulsed unipolar plasma electrolytic oxidation, Philos. Mag. 88 (2008) 795–807. [29] B.D. Ng, I. Annergren, A.M. Soutar, K.A. Khor, A.E.W. Jarfors, Characterisation of a duplex TiO2/CaP coating on Ti6Al4V for hard tissue replacement, Biomaterials 26 (2005) 1087–1095. [30] G.H. Lv, H. Chen, W.C. Gu, L. Li, E.W. Niu, X.H. Zhang, S.Z. Yang, Effects of current frequency on the structural characteristics and corrosion property of ceramic coatings formed on magnesium alloy by PEO technology, J. Mater. Process. Technol. 208 (2008) 9–13. [31] K. Shimazaki, V. Mooney, Comparative-study of porous hydroxyapatite and tricalcium phosphate as bone substitute, J. Orthopaed. Res. 3 (1985) 301–310. [32] K.A. Hing, S.M. Best, K.E. Tanner, W. Bonfield, P.A. Revell, Quantification of bone ingrowth within bone-derived porous hydroxyapatite implants of varying density, J. Mater. Sci.:Mater. M 10 (1999) 663–670. [33] Y.Q. Fu, A.W. Batchelor, K.A. Khor, Fretting wear behavior of thermal sprayed hydroxyapatite coating lubricated with bovine albumin, Wear 230 (1999) 98–102. [34] K.G. Budinski, Tribological properties of titanium-alloys, Wear 151 (1991) 203–217. [35] M. Long, H.J. Rack, Titanium alloys in total joint replacement—a materials science perspective, Biomaterials 19 (1998) 1621–1639. [36] B. Koch, J.G.C. Wolke, K. Degroot, X-ray-diffraction studies on plasma-sprayed calcium phosphate-coated implants, J. Biomed. Mater. Res. 24 (1990) 655–667. [37] S.R. Radin, P. Ducheyne, Plasma spraying induced changes of calciumphosphate ceramic characteristics and the effect on invitro stability, J. Mater. Sci.: Mater. M 3 (1992) 33–42. [38] M.J. Filiaggi, R.M. Pilliar, N.A. Coombs, Post-plasma-spraying heat-treatment of the Ha coating/Ti–6Al–4V implant system, J. Biomed. Mater. Res. 27 (1993) 191–198. [39] H.X. Ji, P.M. Marquis, Effect of heat-treatment on the microstructure of plasmasprayed hydroxyapatite coating, Biomaterials 14 (1993) 64–68. [40] X.B. Xiong, X.R. Zeng, C.L. Zou, P. Li, Y.B. Fan, Influence of hydrothermal temperature on hydroxyapatite coating transformed from monetite on HT-C/C composites by induction heating method, Surf. Coat. Technol. 204 (2009) 115–119. [41] O. Blind, L.H. Klein, B. Dailey, L. Jordan, Characterization of hydroxyapatite films obtained by pulsed-laser deposition on Ti and Ti–6AL–4V substrates, Dent. Mater. 21 (2005) 1017–1024.