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Electrodeposition of hydroxyapatite coating on AZ91D magnesium alloy for biomaterial application
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Materials Letters 62 (2008) 3276 – 3279 www.elsevier.com/locate/matlet
Electrodeposition of hydroxyapatite coating on AZ91D magnesium alloy for biomaterial application Y.W. Song ⁎, D.Y. Shan, E.H. Han Environmental Corrosion Center, Institute of Metal Research, The Chinese Academy of Sciences, 62 Wencui Road, Shenyang, 110016, China Received 17 October 2007; accepted 19 February 2008 Available online 26 February 2008
Abstract Magnesium and its alloys are potential biodegradable implant materials due to their attractive biological property. But their poor corrosion resistance may result in the sudden failure of the implants. The bioactive hydroxyapatite (HA) coating was electrodeposited on AZ91D magnesium alloy surface to improve its biodegradation performance. The biodegradable behavior of HA coating was investigated by electrochemical tests and immersion tests. The experimental results indicated that the as-deposited coating consisting of dicalcium phosphate dehydrate ((DCPD, CaHPO4 d 2H2O) and β-tricalcium phosphate (β-TCP, Ca3 (PO4)2) was transformed into uniform hydroxyapatite (HA, Ca10 (PO4)6(OH)2) coating after immersion in 1 M NaOH solution for 2 h. The HA coating can obviously slow down the biodegradation rate of AZ91D magnesium alloy in stimulated body fluid (SBF). © 2008 Elsevier B.V. All rights reserved. Keywords: Deposition; Hydroxyapatite; Coating; Biodegradable property; Magnesium alloy
1. Introduction Traditional biomedical implant materials such as stainless steels and Ti alloys play an important role in repairing the damaged bone tissue. If these implants exist in the human body for long time, they will always release toxic elements to impair human body's heath. The application of biodegradable implants can solve this problem. The biodegradable implants can gradually be dissolved, absorbed, consumed or excreted after the bone tissue healing. Current biodegradable implants of polymers have an unsatisfactory mechanical property [1]. In comparison, magnesium and its alloys are potential biodegradable materials due to their attractive biological performances [2–4]: (1) metal magnesium is biodegradable in body fluids by corrosion; (2) Mg2+ is harmless to human body; (3) magnesium can accelerate the growth of new bone tissue; (4) the density, elastic modulus and yield strength of magnesium are closer to the bone tissue than that of the conventional implants. Thus, magnesium and its alloys are superior to any other metallic or polymeric implants at bone repairing or orthopedics. ⁎ Corresponding author. Tel.: +86 24 23915897; fax: +86 24 23894149. E-mail address: [email protected] (Y.W. Song). 0167-577X/$ - see front matter © 2008 Elsevier B.V. All rights reserved. doi:10.1016/j.matlet.2008.02.048
However, magnesium and its alloys are susceptible to suffer attack in chloride containing solutions, e.g. the human body fluid or blood plasma [5]. If the implants being made of magnesium alloys are used to repair the diseased bone tissue, they are possible to lose the mechanical property before the healing of bone tissue due to the rapid corrosion. Recently, some researches have been done to slow down the biodegradation rate of magnesium alloys, including fluoride conversion coating [6], alkali-heat treatment [7] and plasma immersion ion implantation [8]. Besides improving the biodegradation rate of magnesium alloys, the biocompatibility should also be considered. The HA [Ca10 (PO4)6(OH)2] coating can satisfy the dual properties. HA is a major inorganic component of natural bone and can accelerate the bone growth [9,10]. But the mechanical strength of HA is too poor to be used in load-bearing applications. Therefore, HA coating was deposited on the surface of metallic implants to improve the biocompatibility property. Until now, these researches about HA coating have mainly focused on the implants of titanium and titanium alloys. The methods to deposit HA coating on titanium alloys surface included plasma spraying, sputtering, pulsed laser-deposition, sol–gel, electrophoresis and electrodeposition, etc. [11]. Among these methods, electrodeposition HA coating, an inexpensive and simple process, can be carried out at room temperature, and the thickness and
Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
Fig. 1. Surface morphology and structure of the as-deposited coating.
Fig. 2. Surface morphology and structure of HA coating.
Fig. 3. Potentiodynamic curves and EIS plots of HA coating and AZ91D in SBF.
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Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
chemical composition of the HA coating can also be controlled by adjusting the electrodeposition conditions [12]. Thus, the aim of this paper is to electrodeposit HA coating on magnesium alloys surface to improve the biodegradation property. 2. Experimental The substrate material was AZ91D magnesium alloy with a size of 30 mm × 25 mm × 5 mm. The samples surface was ground with 1000 grit SiC paper to ensure the same surface roughness. The electrolyte solution contains 0.1 M Ca (NO3)2, 0.06 M NH4H2PO4, H2O2 10 ml/l and pH 4.3. Electrodepositon was carried out at a stable cathodic potential of 4 V for 2 h at room temperature. Then the as-deposited coating was immersed in 1 M NaOH solution for 2 h at 80 °C. The surface morphologies were observed with Philips-XL30 SEM equipped with EDS. The coating structures were analyzed using Philips PW1700 XRD. The biodegradable property was evaluated by electrochemical tests and immersion tests in the stimulated body fluid (SBF). SBF is composed of [2]8.8 g/l NaCl, 0.4 g/l KCl, 0.14 g/l CaCl2, 0.35 g/ l NaHCO3, 1.0 g/l C6H6O6 (glucose), 0.2 g/l MgSO4d 7H2O, 0.1 g/l KH2PO4d H2O, 0.06 g/l Na2HPO4d 7H2O, pH 7.4, temperature 37 °C. Electrochemical tests were carried out using a classical three electrodes cell with platinum as counter electrode, saturated calomel electrode SCE (+0.242 V vs. SHE) as reference electrode and the samples as working electrode. The potentiodynamic polarization curves were obtained using a EG&G potentiostat model 273 at a constant voltage scan rate of 0.3 mV/s. Electrochemical impedance spectroscopy (EIS) measurements were performed using a model 5210 lock in amplifier coupled with potentiostat model 273. The scan frequency ranged from 100 kHz to 10 mHz, and the perturbation amplitude was 5 mV. Immersion tests were done in SBF for 48 h, and then the surface morphologies were observed with SEM. 3. Results and discussion Fig. 1 shows the surface morphology and structure of the asdeposited coating. The coating surface exhibited two kinds of different
crystal characteristics. One appeared regular flake-like structure diverging from centre toward periphery (marked in the white circle in Fig. 1). The others appeared irregular flake-like structure with different dimension. The XRD result indicated that the as-deposited coating mainly consisted of dicalcium phosphate dehydrate (DCPD, CaHPO4d 2H2O) and β-tricalcium phosphate [β-TCP, Ca3 (PO4)2]. The electrodeposition reactions on the surface of AZ91D are suggested as follows [13]. Stage I: Reduction reaction of H2PO−4 and HPO2− 4 2H2 PO−4 þ 2e− →2HPO2− 4 þ H2 ↑
ð1Þ
− 3− 2HPO2− 4 þ 2e →2PO4 þ H2 ↑ 2+
HPO2− 4
ð2Þ PO3− 4
and to form CaHPO4 · Stage II: Ca reacted with 2H2O (DCPD) and Ca3 (PO4)2 (β-TCP), respectively. Ca2þ þ HPO2− 4 þ 2H2 O→CaHPO4 d 2H2 O
ð3Þ
3Ca2þ þ 2PO3− 4 →Ca3 ðPO4 Þ2
ð4Þ
DCPD and β-TCP were the precursors of HA. HAwas the best stable calcium phosphate ceramic in alkaline solution. The as-deposited coating consisting of DCPD and β-TCP was transformed into HA after immersed in 1 M NaOH solution for 2 h Fig. 2 shows the surface morphology and structure of HA coating. The HA coating was not compact and displayed uniform flake-like morphology. The loose HA coating was helpful for the bone tissue to infiltrate into the implants then to accelerate the healing of the damaged bone [14]. Fig. 3 shows the potentiodynamic curves and EIS plots of AZ91D magnesium alloy with and without HA coating in SBF. From the potentiodynamic curves, it was found that the corrosion current density for AZ91D substrate increased quickly at the beginning of anodic side. Then the diffusion-controlled anodic current behavior was observed at the end of the curves due to the fast corrosion rate [15]. It indicated that magnesium alloy substrate suffered severe attack in SBF. After AZ91D was protected with HA coating, the anodic current density increased slowly from free corrosion potential (Ecorr) to approximate − 1.36 V vs. SCE. Then the corrosion current density increased quickly from −1.36 to −1.30 V vs. SCE, indicting the break down of HA coating. This result showed that the HA coating can prevent magnesium alloy from corroding when the potential was more negative than − 1.36 V vs. SCE. Then the magnesium alloy substrate suffered severe attack with
Fig. 4. Corrosion morphologies of AZ91D and HA coating immersed in SFB solution for 48 h (a) AZ91D; (b) HA coating.
Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
increasing of anodic potential. Additional, the free corrosion current (icorr) for the HA coating was 3.65E− 5 A/cm2, which was almost ten times lower than that of AZ91D substrate (2.97E− 4 A/ cm2). Because the cathodic hydrogen evolution reaction was reduced after AZ91D substrate covering with HA coating, the free corrosion potential (Ecorr) of HA coating was more negative than that of the substrate material [15]. If there were pores or crevices in the coating, it can not accelerate the corrosion of magnesium alloys as the metal coating of nickel [16]. Thus, it implied that the HA coating can improve the biodegradable property of magnesium alloys in SBF. According to the EIS plots, obvious change can be found due to the presence of HA coating. The magnesium alloy substrate exhibited two capacitance loops at high frequency and low frequency, respectively. The high frequency capacitance loop described the characteristics of electric double layer. The low frequency loop was related to the adsorption of corrosion products on the AZ91D surface [17]. The plot for the HA coating only contained one capacitance loop, which implied that the coating was undamaged. Additional, the capacitance loop diameter of the HA coating was bigger than that of the magnesium alloy substrate. Thus, the HA coating can reduce the biodegradation rate of magnesium alloy in the SBF. The biodegradable behaviors were studied by immersion tests in SBF for 48 h as shown in Fig. 4. The magnesium alloy substrate suffered serious attack due to the presence of Cl− in SBF. Plenty of white corrosion products were formed and many cracks can also be observed. The sample with the protection of HA coating only suffered attack to some extent. The HA coating surface was not as uniform as without corrosion. Parts of the flake-like HA coating were dissolved into SBF. But the corrosion didn't penetrate the coating. The HA coating can continue to protect the magnesium alloys substrate from corroding for longer time.
4. Conclusions The bioactive hydroxyapatite (HA) coating was obtained by electrodepostion method to improve the biodegradation behaviors of magnesium alloys in human body environment. The asdeposited coating consisted of dicalcium phosphate dehydrate (DCPD) and β-tricalcium phosphate (β-TCP). After immersion in 1 M NaOH solution for 2 h, the as-deposited coating was
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transformed into uniform hydroxyapatite (HA). The hydroxyapatite coating can obviously improve the biodegradation rate of magnesium alloys in stimulated body fluid (SBF) based on the electrochemical results. The corrosion morphologies also indicated that the hydroxyapatite coating can provide enough protection to the magnesium alloy substrate. Acknowledgements This work was supported by National Key Basic Research Program (No.2007CB613705) and National Key Technology R&D Program (2006BAE04B05-2). References [1] C.L. Liu, Y.C. Xin, G.Y. Tang, P.K. Tang, Mater. Sci. Eng., A 456 (2007) 350–357. [2] G.L. Song, Corros. Sci. 49 (2007) 1696–1701. [3] F. Witte, V. Kaese, H. Haferkamp, E. Switzer, Biomaterials 26 (2005) 3557–3563. [4] F. Witte, J. Fischer, J. Nellesen, Biomaterials 27 (2006) 1013–1018. [5] G.L. Song, Adv. Eng. Mater. 7 (2005) 563–586. [6] K.Y. Chiu, M.H. Wong, F.T. Wong, H.C. Wong, Surf. Coat. Technol. 202 (2007) 590–598. [7] L.C. Li, J.C. Li, Y. Li, Surf. Coat. Technol. 185 (2007) 92–98. [8] C.L. Liu, Y.C. Liu, X.B. Tian, P.K. Tian, Thin Solid Films 516 (2007) 422–427. [9] K.J.L. Burg, S. Porter, J.F. Kellam, Biomaterials 21 (2000) 2347–2359. [10] R. Narayanan, S.K. Seshadri, T.Y. Kwon, K.H. Kim, Scripta Mater. 56 (2007) 229–232. [11] E.L. Zhang, K. Yang, Trans. Nonferrous Met. Soc. China 15 (2005) 1199–1205. [12] R. Narayanan, S. Dutta, S.K. Seshadri, Surf. Coat. Technol. 200 (2006) 4720–4730. [13] M.C. Seshadri, S.K. Yen, Mater. Sci. Eng. C 20 (2002) 153–160. [14] J.H. Park, Y.K. Lee, K.M. Kim, Surf. Coat. Technol. 195 (2005) 252–257. [15] C.N. Cao, Principle of corrosion electrochemistry, Chemistry industry press, Beijing, 2004. [16] Y.W. Song, D.Y. Shan, E.H. Han, Electrochim. Acta 53 (2008) 2135–2143. [17] C.N. Cao, J.Q. Zhang, Introduction of Electrochemical Impedance Spectroscopy, Science Press, Beijing, 2002.
Materials Letters 62 (2008) 3276 – 3279 www.elsevier.com/locate/matlet
Electrodeposition of hydroxyapatite coating on AZ91D magnesium alloy for biomaterial application Y.W. Song ⁎, D.Y. Shan, E.H. Han Environmental Corrosion Center, Institute of Metal Research, The Chinese Academy of Sciences, 62 Wencui Road, Shenyang, 110016, China Received 17 October 2007; accepted 19 February 2008 Available online 26 February 2008
Abstract Magnesium and its alloys are potential biodegradable implant materials due to their attractive biological property. But their poor corrosion resistance may result in the sudden failure of the implants. The bioactive hydroxyapatite (HA) coating was electrodeposited on AZ91D magnesium alloy surface to improve its biodegradation performance. The biodegradable behavior of HA coating was investigated by electrochemical tests and immersion tests. The experimental results indicated that the as-deposited coating consisting of dicalcium phosphate dehydrate ((DCPD, CaHPO4 d 2H2O) and β-tricalcium phosphate (β-TCP, Ca3 (PO4)2) was transformed into uniform hydroxyapatite (HA, Ca10 (PO4)6(OH)2) coating after immersion in 1 M NaOH solution for 2 h. The HA coating can obviously slow down the biodegradation rate of AZ91D magnesium alloy in stimulated body fluid (SBF). © 2008 Elsevier B.V. All rights reserved. Keywords: Deposition; Hydroxyapatite; Coating; Biodegradable property; Magnesium alloy
1. Introduction Traditional biomedical implant materials such as stainless steels and Ti alloys play an important role in repairing the damaged bone tissue. If these implants exist in the human body for long time, they will always release toxic elements to impair human body's heath. The application of biodegradable implants can solve this problem. The biodegradable implants can gradually be dissolved, absorbed, consumed or excreted after the bone tissue healing. Current biodegradable implants of polymers have an unsatisfactory mechanical property [1]. In comparison, magnesium and its alloys are potential biodegradable materials due to their attractive biological performances [2–4]: (1) metal magnesium is biodegradable in body fluids by corrosion; (2) Mg2+ is harmless to human body; (3) magnesium can accelerate the growth of new bone tissue; (4) the density, elastic modulus and yield strength of magnesium are closer to the bone tissue than that of the conventional implants. Thus, magnesium and its alloys are superior to any other metallic or polymeric implants at bone repairing or orthopedics. ⁎ Corresponding author. Tel.: +86 24 23915897; fax: +86 24 23894149. E-mail address: [email protected] (Y.W. Song). 0167-577X/$ - see front matter © 2008 Elsevier B.V. All rights reserved. doi:10.1016/j.matlet.2008.02.048
However, magnesium and its alloys are susceptible to suffer attack in chloride containing solutions, e.g. the human body fluid or blood plasma [5]. If the implants being made of magnesium alloys are used to repair the diseased bone tissue, they are possible to lose the mechanical property before the healing of bone tissue due to the rapid corrosion. Recently, some researches have been done to slow down the biodegradation rate of magnesium alloys, including fluoride conversion coating [6], alkali-heat treatment [7] and plasma immersion ion implantation [8]. Besides improving the biodegradation rate of magnesium alloys, the biocompatibility should also be considered. The HA [Ca10 (PO4)6(OH)2] coating can satisfy the dual properties. HA is a major inorganic component of natural bone and can accelerate the bone growth [9,10]. But the mechanical strength of HA is too poor to be used in load-bearing applications. Therefore, HA coating was deposited on the surface of metallic implants to improve the biocompatibility property. Until now, these researches about HA coating have mainly focused on the implants of titanium and titanium alloys. The methods to deposit HA coating on titanium alloys surface included plasma spraying, sputtering, pulsed laser-deposition, sol–gel, electrophoresis and electrodeposition, etc. [11]. Among these methods, electrodeposition HA coating, an inexpensive and simple process, can be carried out at room temperature, and the thickness and
Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
Fig. 1. Surface morphology and structure of the as-deposited coating.
Fig. 2. Surface morphology and structure of HA coating.
Fig. 3. Potentiodynamic curves and EIS plots of HA coating and AZ91D in SBF.
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Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
chemical composition of the HA coating can also be controlled by adjusting the electrodeposition conditions [12]. Thus, the aim of this paper is to electrodeposit HA coating on magnesium alloys surface to improve the biodegradation property. 2. Experimental The substrate material was AZ91D magnesium alloy with a size of 30 mm × 25 mm × 5 mm. The samples surface was ground with 1000 grit SiC paper to ensure the same surface roughness. The electrolyte solution contains 0.1 M Ca (NO3)2, 0.06 M NH4H2PO4, H2O2 10 ml/l and pH 4.3. Electrodepositon was carried out at a stable cathodic potential of 4 V for 2 h at room temperature. Then the as-deposited coating was immersed in 1 M NaOH solution for 2 h at 80 °C. The surface morphologies were observed with Philips-XL30 SEM equipped with EDS. The coating structures were analyzed using Philips PW1700 XRD. The biodegradable property was evaluated by electrochemical tests and immersion tests in the stimulated body fluid (SBF). SBF is composed of [2]8.8 g/l NaCl, 0.4 g/l KCl, 0.14 g/l CaCl2, 0.35 g/ l NaHCO3, 1.0 g/l C6H6O6 (glucose), 0.2 g/l MgSO4d 7H2O, 0.1 g/l KH2PO4d H2O, 0.06 g/l Na2HPO4d 7H2O, pH 7.4, temperature 37 °C. Electrochemical tests were carried out using a classical three electrodes cell with platinum as counter electrode, saturated calomel electrode SCE (+0.242 V vs. SHE) as reference electrode and the samples as working electrode. The potentiodynamic polarization curves were obtained using a EG&G potentiostat model 273 at a constant voltage scan rate of 0.3 mV/s. Electrochemical impedance spectroscopy (EIS) measurements were performed using a model 5210 lock in amplifier coupled with potentiostat model 273. The scan frequency ranged from 100 kHz to 10 mHz, and the perturbation amplitude was 5 mV. Immersion tests were done in SBF for 48 h, and then the surface morphologies were observed with SEM. 3. Results and discussion Fig. 1 shows the surface morphology and structure of the asdeposited coating. The coating surface exhibited two kinds of different
crystal characteristics. One appeared regular flake-like structure diverging from centre toward periphery (marked in the white circle in Fig. 1). The others appeared irregular flake-like structure with different dimension. The XRD result indicated that the as-deposited coating mainly consisted of dicalcium phosphate dehydrate (DCPD, CaHPO4d 2H2O) and β-tricalcium phosphate [β-TCP, Ca3 (PO4)2]. The electrodeposition reactions on the surface of AZ91D are suggested as follows [13]. Stage I: Reduction reaction of H2PO−4 and HPO2− 4 2H2 PO−4 þ 2e− →2HPO2− 4 þ H2 ↑
ð1Þ
− 3− 2HPO2− 4 þ 2e →2PO4 þ H2 ↑ 2+
HPO2− 4
ð2Þ PO3− 4
and to form CaHPO4 · Stage II: Ca reacted with 2H2O (DCPD) and Ca3 (PO4)2 (β-TCP), respectively. Ca2þ þ HPO2− 4 þ 2H2 O→CaHPO4 d 2H2 O
ð3Þ
3Ca2þ þ 2PO3− 4 →Ca3 ðPO4 Þ2
ð4Þ
DCPD and β-TCP were the precursors of HA. HAwas the best stable calcium phosphate ceramic in alkaline solution. The as-deposited coating consisting of DCPD and β-TCP was transformed into HA after immersed in 1 M NaOH solution for 2 h Fig. 2 shows the surface morphology and structure of HA coating. The HA coating was not compact and displayed uniform flake-like morphology. The loose HA coating was helpful for the bone tissue to infiltrate into the implants then to accelerate the healing of the damaged bone [14]. Fig. 3 shows the potentiodynamic curves and EIS plots of AZ91D magnesium alloy with and without HA coating in SBF. From the potentiodynamic curves, it was found that the corrosion current density for AZ91D substrate increased quickly at the beginning of anodic side. Then the diffusion-controlled anodic current behavior was observed at the end of the curves due to the fast corrosion rate [15]. It indicated that magnesium alloy substrate suffered severe attack in SBF. After AZ91D was protected with HA coating, the anodic current density increased slowly from free corrosion potential (Ecorr) to approximate − 1.36 V vs. SCE. Then the corrosion current density increased quickly from −1.36 to −1.30 V vs. SCE, indicting the break down of HA coating. This result showed that the HA coating can prevent magnesium alloy from corroding when the potential was more negative than − 1.36 V vs. SCE. Then the magnesium alloy substrate suffered severe attack with
Fig. 4. Corrosion morphologies of AZ91D and HA coating immersed in SFB solution for 48 h (a) AZ91D; (b) HA coating.
Y.W. Song et al. / Materials Letters 62 (2008) 3276–3279
increasing of anodic potential. Additional, the free corrosion current (icorr) for the HA coating was 3.65E− 5 A/cm2, which was almost ten times lower than that of AZ91D substrate (2.97E− 4 A/ cm2). Because the cathodic hydrogen evolution reaction was reduced after AZ91D substrate covering with HA coating, the free corrosion potential (Ecorr) of HA coating was more negative than that of the substrate material [15]. If there were pores or crevices in the coating, it can not accelerate the corrosion of magnesium alloys as the metal coating of nickel [16]. Thus, it implied that the HA coating can improve the biodegradable property of magnesium alloys in SBF. According to the EIS plots, obvious change can be found due to the presence of HA coating. The magnesium alloy substrate exhibited two capacitance loops at high frequency and low frequency, respectively. The high frequency capacitance loop described the characteristics of electric double layer. The low frequency loop was related to the adsorption of corrosion products on the AZ91D surface [17]. The plot for the HA coating only contained one capacitance loop, which implied that the coating was undamaged. Additional, the capacitance loop diameter of the HA coating was bigger than that of the magnesium alloy substrate. Thus, the HA coating can reduce the biodegradation rate of magnesium alloy in the SBF. The biodegradable behaviors were studied by immersion tests in SBF for 48 h as shown in Fig. 4. The magnesium alloy substrate suffered serious attack due to the presence of Cl− in SBF. Plenty of white corrosion products were formed and many cracks can also be observed. The sample with the protection of HA coating only suffered attack to some extent. The HA coating surface was not as uniform as without corrosion. Parts of the flake-like HA coating were dissolved into SBF. But the corrosion didn't penetrate the coating. The HA coating can continue to protect the magnesium alloys substrate from corroding for longer time.
4. Conclusions The bioactive hydroxyapatite (HA) coating was obtained by electrodepostion method to improve the biodegradation behaviors of magnesium alloys in human body environment. The asdeposited coating consisted of dicalcium phosphate dehydrate (DCPD) and β-tricalcium phosphate (β-TCP). After immersion in 1 M NaOH solution for 2 h, the as-deposited coating was
3279
transformed into uniform hydroxyapatite (HA). The hydroxyapatite coating can obviously improve the biodegradation rate of magnesium alloys in stimulated body fluid (SBF) based on the electrochemical results. The corrosion morphologies also indicated that the hydroxyapatite coating can provide enough protection to the magnesium alloy substrate. Acknowledgements This work was supported by National Key Basic Research Program (No.2007CB613705) and National Key Technology R&D Program (2006BAE04B05-2). References [1] C.L. Liu, Y.C. Xin, G.Y. Tang, P.K. Tang, Mater. Sci. Eng., A 456 (2007) 350–357. [2] G.L. Song, Corros. Sci. 49 (2007) 1696–1701. [3] F. Witte, V. Kaese, H. Haferkamp, E. Switzer, Biomaterials 26 (2005) 3557–3563. [4] F. Witte, J. Fischer, J. Nellesen, Biomaterials 27 (2006) 1013–1018. [5] G.L. Song, Adv. Eng. Mater. 7 (2005) 563–586. [6] K.Y. Chiu, M.H. Wong, F.T. Wong, H.C. Wong, Surf. Coat. Technol. 202 (2007) 590–598. [7] L.C. Li, J.C. Li, Y. Li, Surf. Coat. Technol. 185 (2007) 92–98. [8] C.L. Liu, Y.C. Liu, X.B. Tian, P.K. Tian, Thin Solid Films 516 (2007) 422–427. [9] K.J.L. Burg, S. Porter, J.F. Kellam, Biomaterials 21 (2000) 2347–2359. [10] R. Narayanan, S.K. Seshadri, T.Y. Kwon, K.H. Kim, Scripta Mater. 56 (2007) 229–232. [11] E.L. Zhang, K. Yang, Trans. Nonferrous Met. Soc. China 15 (2005) 1199–1205. [12] R. Narayanan, S. Dutta, S.K. Seshadri, Surf. Coat. Technol. 200 (2006) 4720–4730. [13] M.C. Seshadri, S.K. Yen, Mater. Sci. Eng. C 20 (2002) 153–160. [14] J.H. Park, Y.K. Lee, K.M. Kim, Surf. Coat. Technol. 195 (2005) 252–257. [15] C.N. Cao, Principle of corrosion electrochemistry, Chemistry industry press, Beijing, 2004. [16] Y.W. Song, D.Y. Shan, E.H. Han, Electrochim. Acta 53 (2008) 2135–2143. [17] C.N. Cao, J.Q. Zhang, Introduction of Electrochemical Impedance Spectroscopy, Science Press, Beijing, 2002.