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Plasma-sprayed diopside coatings for biomedical applications
Surface & Coatings Technology 185 (2004) 340 – 345 www.elsevier.com/locate/surfcoat
Plasma-sprayed diopside coatings for biomedical applications Weichang Xue *, Xuanyong Liu, Xuebin Zheng, Chuanxian Ding Shanghai Institute of Ceramics, Chinese Academy of Science, 1295 Dingxi Road, Shanghai 200050, PR China Received 14 July 2003; accepted in revised form 23 December 2003 Available online 30 April 2004
Abstract Diopside coatings have been sprayed onto Ti – 6Al – 4V substrates using an atmospheric plasma spray system. The phase composition and microstructure of the powders and coatings were examined by scanning electron microscopy and X-ray diffraction. The thermal expansion coefficient of diopside coating was measured by a dilatometer in the temperature range of 20 – 600 jC, which adapted to that of titanium alloy. The bond strength of coating was approximately 32.5 MPa, which is higher than that of plasma sprayed HA coating on titanium alloys substrates. The density, porosity, roughness and Young’s moduli of diopside coatings were also measured. The bioactivity of diopside coating was evaluated in vitro by soaking it in simulated body fluid. After 15 days soaking in simulated body fluid, an apatite layer was formed on the surface of the coating. The results are obtained indicated that plasma sprayed diopside coating and may be a candidate for bone and dental implant. D 2004 Elsevier B.V. All rights reserved. Keywords: Diopside coating; Plasma spraying; Thermal expansion coefficient apatite; Simulated body fluid
1. Introduction Hydroxyapatite (HA) coatings deposited on the surface of titanium alloy (Ti– 6Al –4V) implants by plasma spraying are being widely used in orthopedics and dentistry [1]. Such implants combine the strength, ductility and ease of fabrication of titanium alloy with the increased biocompatibility associated with HA. However, some metastable and amorphous phases appear in the HA coating during the plasma spraying process, which result in the low crystallinity of HA coating [2,3]. Long-term animal studies and clinical trials of load-bearing dental and orthopedic prostheses showed that HA coating degrade with time, depending upon the degree of crystallinity of the HA layer [4]. Another crucial problem pertinent to the plasma sprayed HA coating is the poor bonding strength to metal substrate due to the thermal expansion coefficient mismatch at the Ti/HA interface. The thermal expansion coefficient of titanium was approximately 8– 1010 6/jC, which amounts to only 60%
* Corresponding author. Tel.: +86-21-52412990; fax: +86-2152413903. E-mail address: [email protected] (W. Xue). 0257-8972/$ - see front matter D 2004 Elsevier B.V. All rights reserved. doi:10.1016/j.surfcoat.2003.12.018
of that of HA. The mismatch of the thermal expansion coefficient between the substrate and the coating resulted in higher tensile stresses and microcracks at the interface, which also caused a decrease in the bond strength of the coating. It has a tendency to lead to the peeling and also the fatigue failure under tensile loading [5]. To increase bonding between HA coating and Ti substrates, an intermediate layer [6] or a graded coating [7] had been proposed. Diopside (CaMgSi2O6), one of the pyroxene minerals, is a raw material mainly used for traditional ceramics. The thermal expansion coefficient of diopside is similar to that of titanium alloy [8], which was also suitable for the bonding between the substrates and coatings. The bioactivity of diopside has been reported in some studies. Nanami [9] reported apatite to be formed on diopside ceramic surface by soaking in simulated body fluid. It also has been shown that diopside has no cytotoxicity against the fibroblast-like cell line L929 [10] and the osteogenic cell line MC3T3-E1 [11]. It has been demonstrated that diopside has the ability to bond directly to the bone [10]. Therefore, an interest has been taken in diopside as a plasma-sprayed coating for applications in which good mechanical properties and bioactivity are desired. In the present work, diopside coatings on Ti –6Al – 4V substrate were prepared by
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345 Table 1 Plasma spraying parameters Argon plasma gas flow rate (slpm) Hydrogen plasma gas flow rate (slpm) Spray distance (mm) Argon powder carrier gas (slpm) Current (A) Voltage (V) Powder feed rate (g/min)
40 10 120 3.5 600 70 15
atmospheric plasma spraying. The microstructure, phase composition, thermo-physical and some mechanical properties of the coatings were measured. The bioactivity of coating also was evaluated by soaking in simulated body fluid (SBF).
2. Materials and methods 2.1. Substrate preparation and plasma spraying Substrates of Ti– 6Al– 4V with dimensions 20104 mm were grit-blasted and cleaned with acetone before plasma spraying. Commercially available diopside (CaMgSi2O6) powder was used for as-received powder and was sprayed onto substrates by an atmosphere plasma spray system (APS-2000, Sulzer Metco, Switzerland), using the parameters shown in Table 1. Coatings of thickness ranging from 200 to 300 Am were obtained. 2.2. Coating characterization The phase content of the powders and the as-sprayed coatings were examined by an X-ray diffractometer (D/ max 2550v, Japan). The surface morphology and crosssection microstructure of the coating were observed using a scanning electron microscope with electron probe X-ray microanalysis (EPMA-8705QH2, Japan). The tensile bond strength between coating and substrate was measured in accordance with ASTM C-633. For this test, diopside coatings of thickness of approximately 400 Am were sprayed on Ti – 6Al – 4V rods 25.4 mm in diameter. Surface roughness (Ra) was measured by a profilometer (Hommelwerke T8000-C, Germany). For density, Young’s modulus and thermal expansion coefficient measurements, diopside coatings of approximately 3 mm were sprayed onto steel substrates, which were not grit-blasted. After spraying, diopside coatings were removed from substrates. The density and open porosity of coatings were determined by Archimedes’ method. The total porosity of the coatings were estimated by image analysis techniques. Porosity levels were determined by capturing an image of the coating cross-section and differentiating between porosity and bulk coating by grey levels. The thermal expansion coefficient of diopside
341
coating was measured by a dilatometer (402ES-3, Germany) in the temperature range from 20 to 600 jC. A material testing instrument (Instron-5566, UK) was used to measure the Young’s moduli of the coatings by a three-point bending test. All values measured were the mean of five samples. 2.3. Coating bioactivity The bioactivity of diopside coatings was evaluated in vitro by soaking in SBF. After being cleaned with acetone and deionized water in an ultrasonic cleaner, the plates coated with diopside were soaked in SBF (2.5 mM of Ca2+, 1.5 mM of Mg2+, 142.0 mM of Na+, 5.0 mM of K+, 148.5 mM of Cl , 4.2 mM of HCO3 , 1.0 mM of HPO42 , 0.5 mM of SO42 ). The SBF solution was buffered at a pH of 7.40 with 50 mM tri (hydroxymethyl) aminomethane ((CH2OH)3CNH2) and approximately 45 mM hydrochloric acid (HCl) at 37 jC. After soaking for 5 and 15 days, the corresponding surface views of the coatings were observed by scanning electron microscopy (SEM, JEOL JSM-6700F, JAPAN) and energy dispersive spectrometry (EDS, INCA ENERGY, UK).
3. Result and discussion Fig. 1 shows SEM views of the as-received powder (left) and the as-sprayed coating (right). The SEM micrograph of the powder indicated that the particles are mainly angular in shape (Fig. 1a). Fig. 1b shows the surface of a diopside coating that was built up from melted splats with a rough surface, which was suitable to bone implants. The corresponding polished cross-section of the coating is shown in Fig. 2. Pores and microcracks were clearly observed in the coating. The pores formed as a result of poor bonding between adjacent splats, whereas microcracking arose from shrinkage of the splats during the quenching of the melt particles. But, no microcracks was observed at the interface between substrate and coating, which indicated good bonding between substrate and coating. The bond strength of the coating was approximately 32.5 MPa, which was higher than the bond strength of plasma sprayed HA coating reported in some other studies [2,6,12,13]. This can be attributed to the similarity of thermal expansion coefficients between titanium alloy and diopside coating. Fig. 3 shows the results of the measurement of the thermal expansion coefficient of diopside coating compared with those of titanium alloy and HA. The thermal expansion coefficient of diopside coating was approximately 8.4110 6/jC (20 – 600 jC), which was similar to that of titanium alloy (9.4010 6/ jC). The thermal expansion coefficient of HA was approximately 15.2010 6/jC (20 – 600 jC), which was
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Fig. 1. SEM micrographs of (a) the as-received diopside powder and (b) the surface of the as-sprayed diopside coating.
much higher than that of titanium alloy. Because of the mismatch of the thermal expansion coefficient between titanium alloy and HA, tensile stresses may occur during plasma spraying producing microcracks at the interface and causing a decrease in the bond strength. In addition, the thermal expansion coefficient of diopside coating was slightly lower than that of diopside ceramics reported by Hayashi et al. [8]. This can be explained in terms of higher porosity and existence of a glassy phase in the coating. Fig. 4 shows the XRD pattern of diopside powder and coating. As shown in Fig. 4a, major diffraction peaks
Fig. 2. SEM micrograph of the cross-section of diopside coating.
shown in this pattern belong to diopside, and several low intensity peaks originated from cristobalite (SiO2). The XRD pattern of coating indicated the coatings were primarily composed of diopside and an amorphous phase (Fig. 4b). An amorphous phase is often observed in plasma sprayed ceramic coatings. The molten ceramic powder was quenched from a high temperature and solidified at the substrate. From the phase diagram of CaO – MgO – SiO2 [15,16], the akermanite (Ca2MgSi2O7) phase would appear as a result of SiO2 exsolution at high temperature. This was not observed in this study, presumedly due to excess SiO2 existing in the starting powders, as shown by the presence of cristobalite (Fig. 4a).
Fig. 3. Thermal expansion coefficient of diopside coatings compared with that of titanium alloy (Ref. [14]) and hydroxyapatite (Ref. [5]).
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345
Fig. 4. XRD spectra of (a) diopside powder and (b) an as-sprayed diopside coating.
Apart from the bioactivity and bonding strength, other features like density, porosity, surface roughness and Young’s moduli of the coating are essential and crucial for a biomedical coating. Measured coating properties were summarized in Table 2. From Table 2, it can be seen that the density of the diopside coating was measured as 2.52 g/cm3, which is less than the theoretical density of diopside (3.26 g/cm3) [17]. This could be attributed to the presence of some pores and glassy phase in the coatings. The pores provided a mechanical interlock leading to firmer fixation of the material [18]. The porosity of diopside coatings was measured to be approximately 11%. The open porosity of coatings measured by the Archimedes’ method was 5.6%. This means that the volume of the open pores are about half that of the total pores. The size of most pores was less than 10 Am (see Fig. 2). The diopside coating has a rough surface. The surface roughness (Ra) was measured to be 8.30 Am. As shown in Table 2, the Young’s moduli of diopside coating was 38.56 GPa, which is close to that of cortical bone [19] as compared with other bioactive ceramics, such as HA (80 – 110 GPa) and AW glass ceramic (218 GPa) [1].
343
The bioactivity of diopside coating was evaluated in vitro by soaking in SBF. The SEM micrographs of coatings soaked in SBF for 5 and 15 days are shown in Fig. 5. After 5 days soaking in SBF, the isolated granular crystals were observed in the surface of coatings (Fig. 5a). The EDS spectra indicated that the granules were mainly composed of calcium and phosphorus (Fig. 5c). After 15 days soaking in SBF, the surface of coatings was completely covered by ball-like particles (Fig. 5b). At a higher magnification (Fig. 5e), the new-formed layer showed very fine crystallites of approximately 100 nm average size. The EDS quantitative analysis of these fine crystallites gave a Ca/P ratio of approximately 1.62, which was near to the composition of apatite (Fig. 5d). The mechanism of the bone-like apatite formation on the surface of diopside coating in the SBF was similar to that of CaO –SiO2-based glasses [20]. The calcium and magnesium ion released from diopside increases the ionic activity product of the apatite in the surrounding fluid, and hydrated silica on diopside surfaces provided favorable sites for apatite nucleation. Consequently, the apatite was formed on the diopside surface. Once the apatite nuclei were formed, they grew spontaneously by consuming the calcium and phosphate ions from the SBF fluid, since SBF was already supersaturated with respect to apatite. Miake et al. [21] speculated that octacalcium phosphate (OCP) might form at the surface of diopside in SBF or in vivo by epitaxial crystal growth, which changed to HA by a phase transition. Further experimental study is needed for characterizing the diopside coating responsible for apatite formation.
4. Conclusions A new bioceramic coating based on diopside was prepared by plasma spraying. The thermal expansion coefficient of the coating was adapted to that of titanium alloy. The bond strength of the coating was approximately 32.5 MPa, which is higher than that of HA coatings used in orthopedics and dentistry. The Young’s moduli of the coatings was 38.56 GPa, which was close to that of the cortical bone. A fine-grained calcium phosphate (apatite) layer was formed on the surface of coatings soaked in SBF. These Table 2 Thermo-physical and mechanical properties of diopside coatings Property
Value 3
Density (g cm ) Porosity (%) Open porosity (%) Bond strength (MPa) Thermal expansivity coefficient (10 6/jC) (20 – 600 jC) Young’s modulus (GPa) Surface roughness Ra (Am)
2.52F0.02 11.0F1.0 5.65F0.04 32.5F2.8 8.41F0.25 38.56F2.12 8.30F0.31
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Fig. 5. SEM micrographs of diopside coatings after immersion in SBF for (a) 5 days, (b) 15 days, (e) higher magnification of (b), (c) EDS for the granule in (a), (d) EDS for the newly-formed layer in (b).
results suggested that plasma sprayed diopside coating and may be a candidate for bone and dental implant. Acknowledgments This work was supported by Shanghai Science and Technology R&D Fund under grant 02QE14052 and 03JC14074, National Basic Research Fund under grant G1999064701 and National Natural Science Foundation of China 50102008. References [1] L.L. Hench, J. Am. Ceram. Soc. 81 (1998) 1705 – 1728. [2] Y.C. Tsui, C. Doyle, T.W. Clyne, Biomaterials 19 (1998) 2015 – 2029.
[3] Y.C. Tsui, C. Doyle, T.W. Clyne, Biomaterials 19 (1998) 2031 – 2043. [4] L.L. Hench, J. Wilson, An Introduction to Bioceramics, World Scientific, London, UK, 1993, Ch. 12. [5] J. Breme, Y. Zhou, L. Groh, Biomaterials 16 (1995) 239 – 244. [6] D. Lamy, A.C. Pierre, R.B. Heimann, J. Mater. Res. 11 (1996) 680 – 686. [7] E. Park, R.A. Condrate, Mater. Lett. 40 (1999) 228 – 234. [8] K. Hayashi, Y. Nishioka, Y. Okamoto, Nippon. Seram. Kyo. Gak. Ronbunshi. 98 (8) (1990) 801 – 805. [9] T. Nonami, Mat. Res. Soc. Symp. Proc. 252 (1992) 87 – 92. [10] T. Nonami, S. Tsutsumi, J. Mater. Sci. Mater. Med. 10 (1990) 475 – 479. [11] T. Kobayashi, K. Okada, T. Kuroda, K. Sato, J. Biomed. Mater. Res. 37 (1997) 100 – 107. [12] X. Zhen, M. Huang, C. Ding, Biomaterials 21 (2000) 841 – 849. [13] Y. Yang, E. Chang, Biomaterials 22 (2001) 1827 – 1836. [14] Mechanical Engineering Handbook, Mechanical industry press, Beijing, 1996, 4 – 46. [15] H. Takei, T. Miura, J. Cryst. Growth 60 (1982) 453 – 456.
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345 [16] M. Levin Ernest, F. McMurdie Howard, F.P. Hall, Phase Diagrams for Ceramists, The American Ceramic Society, 1956, p. 115. [17] S. Hayashi, K. Okada, N. Otsuka, T. Yano, J. Mater. Sci. Lett. 9 (1990) 382 – 385. [18] W. Suchanek, M. Yoshimura, J. Mater. Res. 13 (1998) 94 – 117.
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[19] L. Hench, J. Am. Ceram. Soc. 74 (1991) 1487 – 1510. [20] T. Kokubo, Thermochimica. Acta 280 – 281 (1996) 479 – 490. [21] Y. Miake, T. Yanagisawa, Y. Yajima, H. Noma, N. Yasui, T. Nonami, J. Dent. Res. 74 (1995) 1756 – 1763.
Plasma-sprayed diopside coatings for biomedical applications Weichang Xue *, Xuanyong Liu, Xuebin Zheng, Chuanxian Ding Shanghai Institute of Ceramics, Chinese Academy of Science, 1295 Dingxi Road, Shanghai 200050, PR China Received 14 July 2003; accepted in revised form 23 December 2003 Available online 30 April 2004
Abstract Diopside coatings have been sprayed onto Ti – 6Al – 4V substrates using an atmospheric plasma spray system. The phase composition and microstructure of the powders and coatings were examined by scanning electron microscopy and X-ray diffraction. The thermal expansion coefficient of diopside coating was measured by a dilatometer in the temperature range of 20 – 600 jC, which adapted to that of titanium alloy. The bond strength of coating was approximately 32.5 MPa, which is higher than that of plasma sprayed HA coating on titanium alloys substrates. The density, porosity, roughness and Young’s moduli of diopside coatings were also measured. The bioactivity of diopside coating was evaluated in vitro by soaking it in simulated body fluid. After 15 days soaking in simulated body fluid, an apatite layer was formed on the surface of the coating. The results are obtained indicated that plasma sprayed diopside coating and may be a candidate for bone and dental implant. D 2004 Elsevier B.V. All rights reserved. Keywords: Diopside coating; Plasma spraying; Thermal expansion coefficient apatite; Simulated body fluid
1. Introduction Hydroxyapatite (HA) coatings deposited on the surface of titanium alloy (Ti– 6Al –4V) implants by plasma spraying are being widely used in orthopedics and dentistry [1]. Such implants combine the strength, ductility and ease of fabrication of titanium alloy with the increased biocompatibility associated with HA. However, some metastable and amorphous phases appear in the HA coating during the plasma spraying process, which result in the low crystallinity of HA coating [2,3]. Long-term animal studies and clinical trials of load-bearing dental and orthopedic prostheses showed that HA coating degrade with time, depending upon the degree of crystallinity of the HA layer [4]. Another crucial problem pertinent to the plasma sprayed HA coating is the poor bonding strength to metal substrate due to the thermal expansion coefficient mismatch at the Ti/HA interface. The thermal expansion coefficient of titanium was approximately 8– 1010 6/jC, which amounts to only 60%
* Corresponding author. Tel.: +86-21-52412990; fax: +86-2152413903. E-mail address: [email protected] (W. Xue). 0257-8972/$ - see front matter D 2004 Elsevier B.V. All rights reserved. doi:10.1016/j.surfcoat.2003.12.018
of that of HA. The mismatch of the thermal expansion coefficient between the substrate and the coating resulted in higher tensile stresses and microcracks at the interface, which also caused a decrease in the bond strength of the coating. It has a tendency to lead to the peeling and also the fatigue failure under tensile loading [5]. To increase bonding between HA coating and Ti substrates, an intermediate layer [6] or a graded coating [7] had been proposed. Diopside (CaMgSi2O6), one of the pyroxene minerals, is a raw material mainly used for traditional ceramics. The thermal expansion coefficient of diopside is similar to that of titanium alloy [8], which was also suitable for the bonding between the substrates and coatings. The bioactivity of diopside has been reported in some studies. Nanami [9] reported apatite to be formed on diopside ceramic surface by soaking in simulated body fluid. It also has been shown that diopside has no cytotoxicity against the fibroblast-like cell line L929 [10] and the osteogenic cell line MC3T3-E1 [11]. It has been demonstrated that diopside has the ability to bond directly to the bone [10]. Therefore, an interest has been taken in diopside as a plasma-sprayed coating for applications in which good mechanical properties and bioactivity are desired. In the present work, diopside coatings on Ti –6Al – 4V substrate were prepared by
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345 Table 1 Plasma spraying parameters Argon plasma gas flow rate (slpm) Hydrogen plasma gas flow rate (slpm) Spray distance (mm) Argon powder carrier gas (slpm) Current (A) Voltage (V) Powder feed rate (g/min)
40 10 120 3.5 600 70 15
atmospheric plasma spraying. The microstructure, phase composition, thermo-physical and some mechanical properties of the coatings were measured. The bioactivity of coating also was evaluated by soaking in simulated body fluid (SBF).
2. Materials and methods 2.1. Substrate preparation and plasma spraying Substrates of Ti– 6Al– 4V with dimensions 20104 mm were grit-blasted and cleaned with acetone before plasma spraying. Commercially available diopside (CaMgSi2O6) powder was used for as-received powder and was sprayed onto substrates by an atmosphere plasma spray system (APS-2000, Sulzer Metco, Switzerland), using the parameters shown in Table 1. Coatings of thickness ranging from 200 to 300 Am were obtained. 2.2. Coating characterization The phase content of the powders and the as-sprayed coatings were examined by an X-ray diffractometer (D/ max 2550v, Japan). The surface morphology and crosssection microstructure of the coating were observed using a scanning electron microscope with electron probe X-ray microanalysis (EPMA-8705QH2, Japan). The tensile bond strength between coating and substrate was measured in accordance with ASTM C-633. For this test, diopside coatings of thickness of approximately 400 Am were sprayed on Ti – 6Al – 4V rods 25.4 mm in diameter. Surface roughness (Ra) was measured by a profilometer (Hommelwerke T8000-C, Germany). For density, Young’s modulus and thermal expansion coefficient measurements, diopside coatings of approximately 3 mm were sprayed onto steel substrates, which were not grit-blasted. After spraying, diopside coatings were removed from substrates. The density and open porosity of coatings were determined by Archimedes’ method. The total porosity of the coatings were estimated by image analysis techniques. Porosity levels were determined by capturing an image of the coating cross-section and differentiating between porosity and bulk coating by grey levels. The thermal expansion coefficient of diopside
341
coating was measured by a dilatometer (402ES-3, Germany) in the temperature range from 20 to 600 jC. A material testing instrument (Instron-5566, UK) was used to measure the Young’s moduli of the coatings by a three-point bending test. All values measured were the mean of five samples. 2.3. Coating bioactivity The bioactivity of diopside coatings was evaluated in vitro by soaking in SBF. After being cleaned with acetone and deionized water in an ultrasonic cleaner, the plates coated with diopside were soaked in SBF (2.5 mM of Ca2+, 1.5 mM of Mg2+, 142.0 mM of Na+, 5.0 mM of K+, 148.5 mM of Cl , 4.2 mM of HCO3 , 1.0 mM of HPO42 , 0.5 mM of SO42 ). The SBF solution was buffered at a pH of 7.40 with 50 mM tri (hydroxymethyl) aminomethane ((CH2OH)3CNH2) and approximately 45 mM hydrochloric acid (HCl) at 37 jC. After soaking for 5 and 15 days, the corresponding surface views of the coatings were observed by scanning electron microscopy (SEM, JEOL JSM-6700F, JAPAN) and energy dispersive spectrometry (EDS, INCA ENERGY, UK).
3. Result and discussion Fig. 1 shows SEM views of the as-received powder (left) and the as-sprayed coating (right). The SEM micrograph of the powder indicated that the particles are mainly angular in shape (Fig. 1a). Fig. 1b shows the surface of a diopside coating that was built up from melted splats with a rough surface, which was suitable to bone implants. The corresponding polished cross-section of the coating is shown in Fig. 2. Pores and microcracks were clearly observed in the coating. The pores formed as a result of poor bonding between adjacent splats, whereas microcracking arose from shrinkage of the splats during the quenching of the melt particles. But, no microcracks was observed at the interface between substrate and coating, which indicated good bonding between substrate and coating. The bond strength of the coating was approximately 32.5 MPa, which was higher than the bond strength of plasma sprayed HA coating reported in some other studies [2,6,12,13]. This can be attributed to the similarity of thermal expansion coefficients between titanium alloy and diopside coating. Fig. 3 shows the results of the measurement of the thermal expansion coefficient of diopside coating compared with those of titanium alloy and HA. The thermal expansion coefficient of diopside coating was approximately 8.4110 6/jC (20 – 600 jC), which was similar to that of titanium alloy (9.4010 6/ jC). The thermal expansion coefficient of HA was approximately 15.2010 6/jC (20 – 600 jC), which was
342
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345
Fig. 1. SEM micrographs of (a) the as-received diopside powder and (b) the surface of the as-sprayed diopside coating.
much higher than that of titanium alloy. Because of the mismatch of the thermal expansion coefficient between titanium alloy and HA, tensile stresses may occur during plasma spraying producing microcracks at the interface and causing a decrease in the bond strength. In addition, the thermal expansion coefficient of diopside coating was slightly lower than that of diopside ceramics reported by Hayashi et al. [8]. This can be explained in terms of higher porosity and existence of a glassy phase in the coating. Fig. 4 shows the XRD pattern of diopside powder and coating. As shown in Fig. 4a, major diffraction peaks
Fig. 2. SEM micrograph of the cross-section of diopside coating.
shown in this pattern belong to diopside, and several low intensity peaks originated from cristobalite (SiO2). The XRD pattern of coating indicated the coatings were primarily composed of diopside and an amorphous phase (Fig. 4b). An amorphous phase is often observed in plasma sprayed ceramic coatings. The molten ceramic powder was quenched from a high temperature and solidified at the substrate. From the phase diagram of CaO – MgO – SiO2 [15,16], the akermanite (Ca2MgSi2O7) phase would appear as a result of SiO2 exsolution at high temperature. This was not observed in this study, presumedly due to excess SiO2 existing in the starting powders, as shown by the presence of cristobalite (Fig. 4a).
Fig. 3. Thermal expansion coefficient of diopside coatings compared with that of titanium alloy (Ref. [14]) and hydroxyapatite (Ref. [5]).
W. Xue et al. / Surface & Coatings Technology 185 (2004) 340–345
Fig. 4. XRD spectra of (a) diopside powder and (b) an as-sprayed diopside coating.
Apart from the bioactivity and bonding strength, other features like density, porosity, surface roughness and Young’s moduli of the coating are essential and crucial for a biomedical coating. Measured coating properties were summarized in Table 2. From Table 2, it can be seen that the density of the diopside coating was measured as 2.52 g/cm3, which is less than the theoretical density of diopside (3.26 g/cm3) [17]. This could be attributed to the presence of some pores and glassy phase in the coatings. The pores provided a mechanical interlock leading to firmer fixation of the material [18]. The porosity of diopside coatings was measured to be approximately 11%. The open porosity of coatings measured by the Archimedes’ method was 5.6%. This means that the volume of the open pores are about half that of the total pores. The size of most pores was less than 10 Am (see Fig. 2). The diopside coating has a rough surface. The surface roughness (Ra) was measured to be 8.30 Am. As shown in Table 2, the Young’s moduli of diopside coating was 38.56 GPa, which is close to that of cortical bone [19] as compared with other bioactive ceramics, such as HA (80 – 110 GPa) and AW glass ceramic (218 GPa) [1].
343
The bioactivity of diopside coating was evaluated in vitro by soaking in SBF. The SEM micrographs of coatings soaked in SBF for 5 and 15 days are shown in Fig. 5. After 5 days soaking in SBF, the isolated granular crystals were observed in the surface of coatings (Fig. 5a). The EDS spectra indicated that the granules were mainly composed of calcium and phosphorus (Fig. 5c). After 15 days soaking in SBF, the surface of coatings was completely covered by ball-like particles (Fig. 5b). At a higher magnification (Fig. 5e), the new-formed layer showed very fine crystallites of approximately 100 nm average size. The EDS quantitative analysis of these fine crystallites gave a Ca/P ratio of approximately 1.62, which was near to the composition of apatite (Fig. 5d). The mechanism of the bone-like apatite formation on the surface of diopside coating in the SBF was similar to that of CaO –SiO2-based glasses [20]. The calcium and magnesium ion released from diopside increases the ionic activity product of the apatite in the surrounding fluid, and hydrated silica on diopside surfaces provided favorable sites for apatite nucleation. Consequently, the apatite was formed on the diopside surface. Once the apatite nuclei were formed, they grew spontaneously by consuming the calcium and phosphate ions from the SBF fluid, since SBF was already supersaturated with respect to apatite. Miake et al. [21] speculated that octacalcium phosphate (OCP) might form at the surface of diopside in SBF or in vivo by epitaxial crystal growth, which changed to HA by a phase transition. Further experimental study is needed for characterizing the diopside coating responsible for apatite formation.
4. Conclusions A new bioceramic coating based on diopside was prepared by plasma spraying. The thermal expansion coefficient of the coating was adapted to that of titanium alloy. The bond strength of the coating was approximately 32.5 MPa, which is higher than that of HA coatings used in orthopedics and dentistry. The Young’s moduli of the coatings was 38.56 GPa, which was close to that of the cortical bone. A fine-grained calcium phosphate (apatite) layer was formed on the surface of coatings soaked in SBF. These Table 2 Thermo-physical and mechanical properties of diopside coatings Property
Value 3
Density (g cm ) Porosity (%) Open porosity (%) Bond strength (MPa) Thermal expansivity coefficient (10 6/jC) (20 – 600 jC) Young’s modulus (GPa) Surface roughness Ra (Am)
2.52F0.02 11.0F1.0 5.65F0.04 32.5F2.8 8.41F0.25 38.56F2.12 8.30F0.31
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Fig. 5. SEM micrographs of diopside coatings after immersion in SBF for (a) 5 days, (b) 15 days, (e) higher magnification of (b), (c) EDS for the granule in (a), (d) EDS for the newly-formed layer in (b).
results suggested that plasma sprayed diopside coating and may be a candidate for bone and dental implant. Acknowledgments This work was supported by Shanghai Science and Technology R&D Fund under grant 02QE14052 and 03JC14074, National Basic Research Fund under grant G1999064701 and National Natural Science Foundation of China 50102008. References [1] L.L. Hench, J. Am. Ceram. Soc. 81 (1998) 1705 – 1728. [2] Y.C. Tsui, C. Doyle, T.W. Clyne, Biomaterials 19 (1998) 2015 – 2029.
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