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The process of electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature
Materials Science and Engineering C 20 (2002) 153 – 160 www.elsevier.com/locate/msec
The process of electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature M.C. Kuo, S.K. Yen* Institute of Materials Engineering, National Chung Hsing University, 250, Kuo Kuang Road, Taichung, 40227, Taiwan, ROC
Abstract Calcium phosphate (CaP) ceramics, especially hydroxyapatite (HA), have received much attention and have been clinically applied in orthopaedics and dentistry due to their excellent biocompatibility. Among several methods for preparing HA coating, electrochemical deposition is a relatively new and possible process. However, documented electrochemical processes were conducted at elevated temperature. In this study, uniform HA coatings have been directly deposited on titanium at room temperature. X-ray diffractometry (XRD) results demonstrated that dicalcium phosphate dihydrate (CaHPO42H2O, DCPD) was the main component of the coating deposited at lower current densities (1 and 5 mA/cm2). HA structure was obtained at current density above 10 mA/cm2 and remained stable after heat treatment at 100 – 600 jC for 1 h. Part of HA phase was transformed into h-TCP and became a biphasic calcium phosphate coating after annealing at 700 jC. Scratch tests showed that HA coating was not scraped off until a shear stress of 106.3 MPa. Coatings deposited at room temperature exhibited stronger adhesion than those at elevated temperature. HA coating revealed a dense inner layer and rough surface morphology which could fulfill the requisition of implant materials and be adequate to the attachment of bone tissue. D 2002 Elsevier Science B.V. All rights reserved. Keywords: Hydroxyapatite; Electrochemical deposition; Ti; Room temperature process
1. Introduction Calcium phosphate (CaP) ceramics have received much attention and have been clinically applied in orthopaedics and dentistry due to their excellent biocompatibility. Among several forms of CaP, synthetic hydroxyapatite [Ca10(PO4)6(OH)2, hereafter HA], a major inorganic component of natural bone, is particularly attractive. HA is recognized as osteoconductive and able to accelerate bone ingrowth and attachment to the surface of implant during the early stages after implantation [1– 3]. Furthermore, the fixation and lifetime of the implant are improved. However, the mechanical strength is too poor to be used in load-bearing applications. Therefore, HA coatings on bioinert metallic prostheses were investigated to maintain biocompatibility and to improve mechanical properties. In the past decade, several methods have been reported to deposit HA onto implant surfaces, for example, plasma spray [3 – 5], rf sputtering [6,7], pulsed laser-deposition [8], sol – gel [9], electrophoretic method [10], and electrochemical deposition [11 –16], etc. Plasma spray has been the
*
Corresponding author. Tel.: +886-4-2840500; fax: +886-4-2857017. E-mail address: [email protected] (S.K. Yen).
commercial method for coating on titanium implant surfaces, but issues remain concerning its use. This includes the presence of calcium phosphate crystalline phases other than HA and amorphous material resulting from the extremely high temperatures used in the plasma spray process [17]. Another issue of greater concern is the poor adhesion of the plasma-sprayed HA coating to titanium substrate due to the delamination of HA coating from the metal implant [18]. Furthermore, plasma spray is difficult in microstructure control and modification, which inherently limits this approach in achieving optimum fixation with the implant. The electrophoretic method is a type of electrodeposition and applies a high potential of 90 V [10] to deposit the coating followed by sintering at high temperatures to improve the adherence. The crystal structures and compositions of the coating are changed after this approach. Calcium phosphate coatings have been deposited on metal substrates by electrochemical method, which is an attractive process because highly irregular objects can be coated relatively quickly at low temperatures. Additionally, the thickness and chemical composition of coatings can be well controlled through adequate conditions of the process. HA, the most interesting form of calcium phosphates was electrochemically deposited in several solutions at elevated temperature [12 – 16], and accompanied with other unstable
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Fig. 1. Cathodic polarization curve of Ti substrate in the electrolyte of 0.042 M Ca(NO3)2 and 0.025 M NH4H2PO4.
calcium phosphate such as dicalcium phosphate dihydrate (CaHPO42H2O, DCPD) [13]. The surface morphology was plate-like and porous in order to provide a relatively large surface area for better chemical interaction with the biological environment. The coatings deposited at room temperature were DCPD in saturated Ca(H2PO4)2 solution [19] and HA salt solution [13]. Ban and Maruno [14] reported
that the coatings deposited at 5 – 37 jC were amorphous in a simulated body fluid, whereas the deposits contained HA at 52 and 62 jC. The main objective of this study was to deposit more crystalline and purer HA coatings at room temperature by the electrochemical method in 0.042 M Ca(NO3)2 and 0.025 M NH4H2PO4 aqueous solution and enhance the adhesion strength on Ti substrate.
Fig. 2. XRD diagrams of specimens deposited at current densities of 1, 5, 10, 15, and 20 mA/cm2 for 30 min, respectively.
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2. Experimental 2.1. Cathodic polarization, deposition, and annealing As-received commercial pure titanium (ASTM F67, grade 2) sheets were cut into disks of diameter 16 mm and thickness 1 mm. After cutting, the specimen were wet ground with SiC abrasive papers, polished to a mirror surface, rinsed with acetone in an ultrasonic bath, and washed with distilled water. The electrolyte used in this study contained 0.042 M Ca(NO 3 ) 2 and 0.025 M NH4H2PO4 and its pH value was 4.11. Cathodic polarization and deposition were carried out with EG&G Model 273A potentiostat/galvanostat. Ti disks were used as the cathode, a Pt plate as the counter-electrode, and a silver/ silver chloride electrode in saturated potassium chloride (0.197 V vs. SHE) as the reference electrode. Cathodic polarization was conducted from open circuit voltage to 3.0 V (vs. Ag/AgCl electrode) at a rate of 0.6 V/h. Coatings were deposited at current densities of 1 –20 mA/ cm2 for duration of 5 –40 min at room temperature (25 jC). After deposition, the specimens were rinsed in distilled water to remove residual electrolyte and dried in air for 24 h. Some specimens deposited at 10 mA/cm2 were annealed at 100 – 700 jC for 1 h. 2.2. XRD, TGA/DTA, SEM/EDS, and scratch tests The crystallography of the as-deposited and annealed specimens was analyzed by X-ray diffractometry (XRD) in a MAC MO3XHF diffractometer with Cu Ka radiation and operated at a tube voltage of 40 kV and a current of 30 mA. The range of 2h was from 10j to 70j with a scanning rate of
155
1j/min. TGA/DTA analysis of the scraped powders from the coatings was operated in ULVAC Sinku-Riko Differential Thermo-gravimetric Analyzer TGD 7000 with temperature range of 25 – 1000 jC at a heating rate of 10 jC/min. The surface morphology and the elemental analysis of the coatings were observed by scanning electron microscopy/energy dispersive spectroscopy (SEM/EDS, JEOL JSM-5400 Japan). The scratch test was conducted to evaluate the coating adhesion. As-deposited specimen was scratched using TEER coating limited (England) Model ST200/F with a pre-load of 1 N, load speed 50 N/min, scratch speed 5 mm/ min, and the maximum load of 40 N.
3. Results and discussion 3.1. Cathodic reactions and crystal structures The cathodic polarization curve of Ti substrate in electrolyte shown in Fig. 1 could be divided into three stages and the related main reactions are suggested as follows. Stage I: Reduction of oxygen (open circuit voltage: 0.4 V) O2 þ 2H2 O þ 4e ! 4OH
ð1Þ
2 Stage II: Reduction of H2PO ( 0.4 to 4 and HPO4 1.6 V)
2 2H2 PO 4 þ 2e ! 2HPO4 þ H2
ð2Þ
3 HPO2 4 þ 2e ! 2PO4 þ H2
ð3Þ
Fig. 3. Weight gain of HA-coated specimens versus deposition time deposited 10 mA/cm2.
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Fig. 4. TGA/DTA analysis of scraped HA powders deposited at 10 mA/cm2.
Stage III: Reduction of water ( 1.6 to 3.0 V) 2H2 O þ 2e ! H2 þ 2OH
ð4Þ
Positive Ca2 + ions migrated to cathodic Ti substrate and could react with PO43 and OH ions formed on Ti surface to synthesize HA. Therefore, current densities of 1 –20 mA/
cm2 were chosen to deposit the coatings, and the corresponding voltages were at Stage III. XRD diagrams of as-deposited coatings are shown in Fig. 2. DCPD was the major and HA the minor structures of the films deposited at a current density of 1 mA/cm2. This result also indicated that reaction (2) was major and reaction (3) was minor at the end of Stage II. When the current density was increased to 5 mA/cm2, DCPD was still the
Fig. 5. XRD diagram of HA-coated specimens deposited at 10 mA/cm2 and then annealed at 200, 400, 600, and 700 jC for 1 h, respectively.
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major structure, but HA increased relatively. However, the diagram illustrated a stronger HA (002) peak at 2h = 25.88j and a broadened HA (211) peak at 2h = 31.77j. When the current density was greater than 10 mA/cm2, the peaks of DCPD dramatically disappeared and HA became the main structure. Therefore, a purer HA coating was obtained by this process. As suggested above, the major reaction at Stage III was reaction (4), the reduction of water, and a lot of OH ions were formed. However, at the beginning of water reduction such as 1– 5 mA/cm2, the concentration of OH was insufficient to convert all HPO42 into PO43 by the following reaction. 3 HPO2 4 þ OH ! PO4 þ H2 O
ð5Þ
Only at a greater current density such as 10 –20 mA/cm2, the concentration of OH was enough to convert HPO42 into PO 43 . Additionally, the broaden peaks of HA revealed the coatings with poor crystallinity, which is similar to nature bone mineral [20,21], and they might be suitable for tissue compatibility. The diffraction peaks were sharper for the deposits obtained at elevated temperature [12]. The weight gain of HA coatings deposited for 5 –40 min is shown in Fig. 3. The weight gain of the coatings increased with increasing duration time for the initial 20 min, reducing slightly after 20 min as since part of the coating was flaked off by the greater amount of H2 bubbles.
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the phase transformation of HA as the following reaction [22], Ca10 ðPO4 Þ6 ðOHÞ2 ! 2Ca3 ðPO4 Þ2 þ Ca2 P2 O7 þ 2CaO þ H2 O
ð6Þ
However, a Ca2P2O7 peak was not found in the XRD diagram. Therefore, another reaction, Ca10 ðPO4 Þ6 ðOHÞ2 ! 3Ca3 ðPO4 Þ2 þ CaO þ H2 O
ð7Þ
should be more feasible. The coating annealed at 700 jC mainly consisted of HA and h-TCP was the attractive biphasic calcium phosphate (BCP) [23]. 3.3. SEM observations and scratch tests SEM photographs of the coatings deposited at 1 and 10 mA/cm2 are shown in Fig. 6(a) and (b), respectively. The coating deposited at 1 mA/cm2 was mainly DCPD and presented plate-like morphology. The HA coating deposited at 10 mA/cm2 illustrated rough particle-like structure on the surface which was different to the results in the literatures [12 –16]. Temperature under deposition and small amount
3.2. TGA/DTA and phase transformation Fig. 4 shows the result of TGA/DTA analysis of scraped HA powders. An endothermic peak and an exothermic peak were detected at 91.1 and 449.0 jC, respectively. A large amount of weight loss was observed at 91.1 jC and a slight weight loss at 725.2 jC. The weight gain at T z 737.5 jC was the baseline of the alumina crucible used in the TGA/ DTA test. XRD diagrams of HA-coated specimens annealed at 200, 400, 600, and 700 jC for 1 h are shown in Fig. 5. The XRD diagram of HA-coated specimen annealed at 200 jC revealed no obvious variation compared with as-deposited specimen. According to TGA/DTA and XRD analysis, the endothermic peak at 91.1 jC was due to the evaporation of water contained in the powders. The increased HA (210) peaks with increasing annealed temperature especially from 400 to 600 jC demonstrated that the exothermic reaction around 449.0 jC was attributed to the crystallization of HA. More crystalline HA coating can be obtained after annealing around 449.0 jC compared with as-deposited HA. h-TCP and CaO found in XRD diagram of HA-coated specimen after annealing at 700 jC elucidated that the weight loss around 725.2 jC on TGA curve was possibly attributed to
Fig. 6. SEM photographs of specimens deposited at (a) 1 mA/cm2 and (b) 10 mA/cm2 for 30 min.
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of DCPD contained in the deposits obtained at elevated temperature were the possible reasons for this difference. The cross-sectional SEM photograph of HA coating deposited at 10 mA/cm2 for 30 min shown in Fig. 7(a) revealed the dense underlayer coating. Fig. 7(b) showed the EDS line-scanning diagram of the straight line AB on Fig. 7(a). The calcium and phosphorous compositions of the coating were uniform and the coating thickness was about 18.6 Am. Ca/P ratios of the coatings deposited at 1 and 10 mA/cm2 were detected by EDS and their values were 1.11 and 1.48, respectively. Ca/P ratio of the coating deposited at 1 mA/ cm2 was slightly greater than the theoretical value for DCPD of 1.00 due to containing a little of HA. The HA coating obtained at 10 mA/cm2 exhibited inferior Ca/P ratio to the
theoretical value of 1.67, thus the HA coating through this process was calcium-deficient HA (CDHA). The friction –load curve of the as-deposited specimen for scratch test and SEM backscattering electron image (BEI) photographs of the scratch trace are shown in Fig. 8(a) and (b), respectively. The HA coating was not scraped off by diamond indenter until the load of 20 N. Part of HA was scraped off when the load was greater than 20 N. All of the coating was scraped off when the load was greater than 23 N where BEI photographs revealed bright titanium signals. With a corresponding friction of 2.85 N at the load of 30 N and 85.94 Am width of the scratch, the shear stress at this stage is about 106.3 MPa, which is three times greater than the loading stress on the hip joint during gait ( < 35 MPa)
Fig. 7. (a) Cross-sectional SEM photograph of HA-coated specimen deposited at 10 mA/cm2 and annealed at 500 jC for 1 h, and (b) EDS line scanning spectrum of line AB in (a).
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Fig. 8. (a) Friction – load curve of HA-coated specimen for scratch test, and (b) SEM/BEI photographs of the scratch trace.
[24], and also greater than the adhesion strength of HA coatings deposited by plasma spray (50 – 60 MPa) [25] and sol – gel methods (20 –30 MPa) [26], respectively.
4. Conclusions Hydroxyapatite (HA) coatings have been deposited on titanium by an electrochemical method at room temperature and presented several differences with respect to other reported works. A purer and more crystalline HA coating formation was formed, based on the greater amount of hydroxyl through the reduction of water and further annealing. The composition of coatings deposited at lower current densities was mainly dicalcium phosphate dihydrate (DCPD). An adherent HA coating synthesized at current densities above 10 mA/cm2 was able to resist a high scratch load of 20 N (i.e. shear stress of 106.3 MPa). The poorly crystalline HA was similar to natural bone mineral. The HA coating derived by electrolytic deposition was a stable phase when annealed at T < 700 jC, and became biphasic calcium phosphate at T z 700 jC.
Acknowledgements The authors are grateful for the support of this research by National Science Council, Republic of China under contract no. 89-2213-E-005-046.
References [1] M. Jarcho, Clin. Orthop. Relat. Res. 157 (1981) 259. [2] S.D. Cook, K.A. Thomas, J.F. Kay, M. Jarcho, Clin. Orthop. Relat. Res. 232 (1988) 225. [3] K. de Groot, R. Geesink, C.P.A.T. Klein, P. Serekian, J. Biomed. Mater. Res. 21 (1987) 1375. [4] J. Weng, Q. Liu, J.G.C. Wolke, X. Zhang, K. de Groot, Biomaterials 18 (1997) 1027. [5] J.H. Chern Lin, M.L. Liu, C.P. Ju, J. Mater. Sci.: Mater. Med. 28 (1994) 1. [6] J.G.C. Wolke, J.P.C.M. van der Waerden, K. de Groot, J.A. Jansen, Biomaterials 18 (1997) 483. [7] S.J. Ding, C.P. Ju, J.H. Chern Lin, J. Biomed. Mater. Res. 44 (1999) 266. [8] C.K. Wang, J.H. Chern Lin, C.P. Ju, H.C. Ong, R.P.H. Chang, Biomaterials 18 (1997) 1331. [9] C.M. Lopatin, V. Pizziconi, T.L. Alford, T. Laursen, Thin Solid Films 326 (1998) 227. [10] P. Ducheyne, S. Radin, M. Heughebaert, J.C. Heughebaert, Biomaterials 11 (1990) 244. [11] M. Gottlander, C.B. Hohansson, A. Wennerberg, T. Albrektsson, S. Ducheyne, P. Ducheyne, Biomaterials 18 (1997) 551. [12] M. Shirkhanzadeh, J. Mater. Sci. Lett. 10 (1991) 1415. [13] T.V. Vijayaraghavan, A. Bensalem, J. Mater. Sci. Lett. 13 (1994) 1782. [14] S. Ban, S. Maruno, Biomaterials 16 (1995) 977. [15] M. Shirkhanzadeh, J. Mater. Sci.: Mater. Med. 9 (1998) 67. [16] M. Shirkhanzadeh, M. Azadegan, Mater. Lett. 15 (1993) 392. [17] J. Huaxia, C.B. Ponton, P.M. Marquis, J. Mater. Sci.: Mater. Med. 3 (1992) 283. [18] S.R. Radin, P. Ducheyne, J. Mater. Sci.: Mater. Med. 3 (1992) 33. [19] J. Redepenning, J.P. McIsaac, Chem. Mater. 2 (1990) 625. [20] R.Z. LeGeros, Prog. Cryst. Growth Charact. 4 (1981) 1. [21] A.S. Posner, Clin. Orthop. Relat. Res. 200 (1985) 87.
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[22] L.M. Rodrigues-Lorenzo, M. Vallet-Regi, J.M.F. Ferreira, Biomaterials 22 (2001) 583. [23] G. Daculsi, Biomaterials 19 (1998) 1473. [24] N. Verdonschot, M. Dalstra, R. Huskes, in: G. Bergmann, F. Graidren, A. Roklmann (Eds.), Implantable Telemetry in Orthopedics, Berlin Free University, Berlin, Germany, 1990, p. 249.
[25] J.H. Chern Lin, H.J. Lin, S.J. Ding, C.P. Ju, Mater. Chem. Phys. 64 (2000) 229. [26] L.-D. Piveteau, B. Gasser, L. Schlapbach, Biomaterials 21 (2000) 2193.
The process of electrochemical deposited hydroxyapatite coatings on biomedical titanium at room temperature M.C. Kuo, S.K. Yen* Institute of Materials Engineering, National Chung Hsing University, 250, Kuo Kuang Road, Taichung, 40227, Taiwan, ROC
Abstract Calcium phosphate (CaP) ceramics, especially hydroxyapatite (HA), have received much attention and have been clinically applied in orthopaedics and dentistry due to their excellent biocompatibility. Among several methods for preparing HA coating, electrochemical deposition is a relatively new and possible process. However, documented electrochemical processes were conducted at elevated temperature. In this study, uniform HA coatings have been directly deposited on titanium at room temperature. X-ray diffractometry (XRD) results demonstrated that dicalcium phosphate dihydrate (CaHPO42H2O, DCPD) was the main component of the coating deposited at lower current densities (1 and 5 mA/cm2). HA structure was obtained at current density above 10 mA/cm2 and remained stable after heat treatment at 100 – 600 jC for 1 h. Part of HA phase was transformed into h-TCP and became a biphasic calcium phosphate coating after annealing at 700 jC. Scratch tests showed that HA coating was not scraped off until a shear stress of 106.3 MPa. Coatings deposited at room temperature exhibited stronger adhesion than those at elevated temperature. HA coating revealed a dense inner layer and rough surface morphology which could fulfill the requisition of implant materials and be adequate to the attachment of bone tissue. D 2002 Elsevier Science B.V. All rights reserved. Keywords: Hydroxyapatite; Electrochemical deposition; Ti; Room temperature process
1. Introduction Calcium phosphate (CaP) ceramics have received much attention and have been clinically applied in orthopaedics and dentistry due to their excellent biocompatibility. Among several forms of CaP, synthetic hydroxyapatite [Ca10(PO4)6(OH)2, hereafter HA], a major inorganic component of natural bone, is particularly attractive. HA is recognized as osteoconductive and able to accelerate bone ingrowth and attachment to the surface of implant during the early stages after implantation [1– 3]. Furthermore, the fixation and lifetime of the implant are improved. However, the mechanical strength is too poor to be used in load-bearing applications. Therefore, HA coatings on bioinert metallic prostheses were investigated to maintain biocompatibility and to improve mechanical properties. In the past decade, several methods have been reported to deposit HA onto implant surfaces, for example, plasma spray [3 – 5], rf sputtering [6,7], pulsed laser-deposition [8], sol – gel [9], electrophoretic method [10], and electrochemical deposition [11 –16], etc. Plasma spray has been the
*
Corresponding author. Tel.: +886-4-2840500; fax: +886-4-2857017. E-mail address: [email protected] (S.K. Yen).
commercial method for coating on titanium implant surfaces, but issues remain concerning its use. This includes the presence of calcium phosphate crystalline phases other than HA and amorphous material resulting from the extremely high temperatures used in the plasma spray process [17]. Another issue of greater concern is the poor adhesion of the plasma-sprayed HA coating to titanium substrate due to the delamination of HA coating from the metal implant [18]. Furthermore, plasma spray is difficult in microstructure control and modification, which inherently limits this approach in achieving optimum fixation with the implant. The electrophoretic method is a type of electrodeposition and applies a high potential of 90 V [10] to deposit the coating followed by sintering at high temperatures to improve the adherence. The crystal structures and compositions of the coating are changed after this approach. Calcium phosphate coatings have been deposited on metal substrates by electrochemical method, which is an attractive process because highly irregular objects can be coated relatively quickly at low temperatures. Additionally, the thickness and chemical composition of coatings can be well controlled through adequate conditions of the process. HA, the most interesting form of calcium phosphates was electrochemically deposited in several solutions at elevated temperature [12 – 16], and accompanied with other unstable
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Fig. 1. Cathodic polarization curve of Ti substrate in the electrolyte of 0.042 M Ca(NO3)2 and 0.025 M NH4H2PO4.
calcium phosphate such as dicalcium phosphate dihydrate (CaHPO42H2O, DCPD) [13]. The surface morphology was plate-like and porous in order to provide a relatively large surface area for better chemical interaction with the biological environment. The coatings deposited at room temperature were DCPD in saturated Ca(H2PO4)2 solution [19] and HA salt solution [13]. Ban and Maruno [14] reported
that the coatings deposited at 5 – 37 jC were amorphous in a simulated body fluid, whereas the deposits contained HA at 52 and 62 jC. The main objective of this study was to deposit more crystalline and purer HA coatings at room temperature by the electrochemical method in 0.042 M Ca(NO3)2 and 0.025 M NH4H2PO4 aqueous solution and enhance the adhesion strength on Ti substrate.
Fig. 2. XRD diagrams of specimens deposited at current densities of 1, 5, 10, 15, and 20 mA/cm2 for 30 min, respectively.
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2. Experimental 2.1. Cathodic polarization, deposition, and annealing As-received commercial pure titanium (ASTM F67, grade 2) sheets were cut into disks of diameter 16 mm and thickness 1 mm. After cutting, the specimen were wet ground with SiC abrasive papers, polished to a mirror surface, rinsed with acetone in an ultrasonic bath, and washed with distilled water. The electrolyte used in this study contained 0.042 M Ca(NO 3 ) 2 and 0.025 M NH4H2PO4 and its pH value was 4.11. Cathodic polarization and deposition were carried out with EG&G Model 273A potentiostat/galvanostat. Ti disks were used as the cathode, a Pt plate as the counter-electrode, and a silver/ silver chloride electrode in saturated potassium chloride (0.197 V vs. SHE) as the reference electrode. Cathodic polarization was conducted from open circuit voltage to 3.0 V (vs. Ag/AgCl electrode) at a rate of 0.6 V/h. Coatings were deposited at current densities of 1 –20 mA/ cm2 for duration of 5 –40 min at room temperature (25 jC). After deposition, the specimens were rinsed in distilled water to remove residual electrolyte and dried in air for 24 h. Some specimens deposited at 10 mA/cm2 were annealed at 100 – 700 jC for 1 h. 2.2. XRD, TGA/DTA, SEM/EDS, and scratch tests The crystallography of the as-deposited and annealed specimens was analyzed by X-ray diffractometry (XRD) in a MAC MO3XHF diffractometer with Cu Ka radiation and operated at a tube voltage of 40 kV and a current of 30 mA. The range of 2h was from 10j to 70j with a scanning rate of
155
1j/min. TGA/DTA analysis of the scraped powders from the coatings was operated in ULVAC Sinku-Riko Differential Thermo-gravimetric Analyzer TGD 7000 with temperature range of 25 – 1000 jC at a heating rate of 10 jC/min. The surface morphology and the elemental analysis of the coatings were observed by scanning electron microscopy/energy dispersive spectroscopy (SEM/EDS, JEOL JSM-5400 Japan). The scratch test was conducted to evaluate the coating adhesion. As-deposited specimen was scratched using TEER coating limited (England) Model ST200/F with a pre-load of 1 N, load speed 50 N/min, scratch speed 5 mm/ min, and the maximum load of 40 N.
3. Results and discussion 3.1. Cathodic reactions and crystal structures The cathodic polarization curve of Ti substrate in electrolyte shown in Fig. 1 could be divided into three stages and the related main reactions are suggested as follows. Stage I: Reduction of oxygen (open circuit voltage: 0.4 V) O2 þ 2H2 O þ 4e ! 4OH
ð1Þ
2 Stage II: Reduction of H2PO ( 0.4 to 4 and HPO4 1.6 V)
2 2H2 PO 4 þ 2e ! 2HPO4 þ H2
ð2Þ
3 HPO2 4 þ 2e ! 2PO4 þ H2
ð3Þ
Fig. 3. Weight gain of HA-coated specimens versus deposition time deposited 10 mA/cm2.
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Fig. 4. TGA/DTA analysis of scraped HA powders deposited at 10 mA/cm2.
Stage III: Reduction of water ( 1.6 to 3.0 V) 2H2 O þ 2e ! H2 þ 2OH
ð4Þ
Positive Ca2 + ions migrated to cathodic Ti substrate and could react with PO43 and OH ions formed on Ti surface to synthesize HA. Therefore, current densities of 1 –20 mA/
cm2 were chosen to deposit the coatings, and the corresponding voltages were at Stage III. XRD diagrams of as-deposited coatings are shown in Fig. 2. DCPD was the major and HA the minor structures of the films deposited at a current density of 1 mA/cm2. This result also indicated that reaction (2) was major and reaction (3) was minor at the end of Stage II. When the current density was increased to 5 mA/cm2, DCPD was still the
Fig. 5. XRD diagram of HA-coated specimens deposited at 10 mA/cm2 and then annealed at 200, 400, 600, and 700 jC for 1 h, respectively.
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major structure, but HA increased relatively. However, the diagram illustrated a stronger HA (002) peak at 2h = 25.88j and a broadened HA (211) peak at 2h = 31.77j. When the current density was greater than 10 mA/cm2, the peaks of DCPD dramatically disappeared and HA became the main structure. Therefore, a purer HA coating was obtained by this process. As suggested above, the major reaction at Stage III was reaction (4), the reduction of water, and a lot of OH ions were formed. However, at the beginning of water reduction such as 1– 5 mA/cm2, the concentration of OH was insufficient to convert all HPO42 into PO43 by the following reaction. 3 HPO2 4 þ OH ! PO4 þ H2 O
ð5Þ
Only at a greater current density such as 10 –20 mA/cm2, the concentration of OH was enough to convert HPO42 into PO 43 . Additionally, the broaden peaks of HA revealed the coatings with poor crystallinity, which is similar to nature bone mineral [20,21], and they might be suitable for tissue compatibility. The diffraction peaks were sharper for the deposits obtained at elevated temperature [12]. The weight gain of HA coatings deposited for 5 –40 min is shown in Fig. 3. The weight gain of the coatings increased with increasing duration time for the initial 20 min, reducing slightly after 20 min as since part of the coating was flaked off by the greater amount of H2 bubbles.
157
the phase transformation of HA as the following reaction [22], Ca10 ðPO4 Þ6 ðOHÞ2 ! 2Ca3 ðPO4 Þ2 þ Ca2 P2 O7 þ 2CaO þ H2 O
ð6Þ
However, a Ca2P2O7 peak was not found in the XRD diagram. Therefore, another reaction, Ca10 ðPO4 Þ6 ðOHÞ2 ! 3Ca3 ðPO4 Þ2 þ CaO þ H2 O
ð7Þ
should be more feasible. The coating annealed at 700 jC mainly consisted of HA and h-TCP was the attractive biphasic calcium phosphate (BCP) [23]. 3.3. SEM observations and scratch tests SEM photographs of the coatings deposited at 1 and 10 mA/cm2 are shown in Fig. 6(a) and (b), respectively. The coating deposited at 1 mA/cm2 was mainly DCPD and presented plate-like morphology. The HA coating deposited at 10 mA/cm2 illustrated rough particle-like structure on the surface which was different to the results in the literatures [12 –16]. Temperature under deposition and small amount
3.2. TGA/DTA and phase transformation Fig. 4 shows the result of TGA/DTA analysis of scraped HA powders. An endothermic peak and an exothermic peak were detected at 91.1 and 449.0 jC, respectively. A large amount of weight loss was observed at 91.1 jC and a slight weight loss at 725.2 jC. The weight gain at T z 737.5 jC was the baseline of the alumina crucible used in the TGA/ DTA test. XRD diagrams of HA-coated specimens annealed at 200, 400, 600, and 700 jC for 1 h are shown in Fig. 5. The XRD diagram of HA-coated specimen annealed at 200 jC revealed no obvious variation compared with as-deposited specimen. According to TGA/DTA and XRD analysis, the endothermic peak at 91.1 jC was due to the evaporation of water contained in the powders. The increased HA (210) peaks with increasing annealed temperature especially from 400 to 600 jC demonstrated that the exothermic reaction around 449.0 jC was attributed to the crystallization of HA. More crystalline HA coating can be obtained after annealing around 449.0 jC compared with as-deposited HA. h-TCP and CaO found in XRD diagram of HA-coated specimen after annealing at 700 jC elucidated that the weight loss around 725.2 jC on TGA curve was possibly attributed to
Fig. 6. SEM photographs of specimens deposited at (a) 1 mA/cm2 and (b) 10 mA/cm2 for 30 min.
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of DCPD contained in the deposits obtained at elevated temperature were the possible reasons for this difference. The cross-sectional SEM photograph of HA coating deposited at 10 mA/cm2 for 30 min shown in Fig. 7(a) revealed the dense underlayer coating. Fig. 7(b) showed the EDS line-scanning diagram of the straight line AB on Fig. 7(a). The calcium and phosphorous compositions of the coating were uniform and the coating thickness was about 18.6 Am. Ca/P ratios of the coatings deposited at 1 and 10 mA/cm2 were detected by EDS and their values were 1.11 and 1.48, respectively. Ca/P ratio of the coating deposited at 1 mA/ cm2 was slightly greater than the theoretical value for DCPD of 1.00 due to containing a little of HA. The HA coating obtained at 10 mA/cm2 exhibited inferior Ca/P ratio to the
theoretical value of 1.67, thus the HA coating through this process was calcium-deficient HA (CDHA). The friction –load curve of the as-deposited specimen for scratch test and SEM backscattering electron image (BEI) photographs of the scratch trace are shown in Fig. 8(a) and (b), respectively. The HA coating was not scraped off by diamond indenter until the load of 20 N. Part of HA was scraped off when the load was greater than 20 N. All of the coating was scraped off when the load was greater than 23 N where BEI photographs revealed bright titanium signals. With a corresponding friction of 2.85 N at the load of 30 N and 85.94 Am width of the scratch, the shear stress at this stage is about 106.3 MPa, which is three times greater than the loading stress on the hip joint during gait ( < 35 MPa)
Fig. 7. (a) Cross-sectional SEM photograph of HA-coated specimen deposited at 10 mA/cm2 and annealed at 500 jC for 1 h, and (b) EDS line scanning spectrum of line AB in (a).
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Fig. 8. (a) Friction – load curve of HA-coated specimen for scratch test, and (b) SEM/BEI photographs of the scratch trace.
[24], and also greater than the adhesion strength of HA coatings deposited by plasma spray (50 – 60 MPa) [25] and sol – gel methods (20 –30 MPa) [26], respectively.
4. Conclusions Hydroxyapatite (HA) coatings have been deposited on titanium by an electrochemical method at room temperature and presented several differences with respect to other reported works. A purer and more crystalline HA coating formation was formed, based on the greater amount of hydroxyl through the reduction of water and further annealing. The composition of coatings deposited at lower current densities was mainly dicalcium phosphate dihydrate (DCPD). An adherent HA coating synthesized at current densities above 10 mA/cm2 was able to resist a high scratch load of 20 N (i.e. shear stress of 106.3 MPa). The poorly crystalline HA was similar to natural bone mineral. The HA coating derived by electrolytic deposition was a stable phase when annealed at T < 700 jC, and became biphasic calcium phosphate at T z 700 jC.
Acknowledgements The authors are grateful for the support of this research by National Science Council, Republic of China under contract no. 89-2213-E-005-046.
References [1] M. Jarcho, Clin. Orthop. Relat. Res. 157 (1981) 259. [2] S.D. Cook, K.A. Thomas, J.F. Kay, M. Jarcho, Clin. Orthop. Relat. Res. 232 (1988) 225. [3] K. de Groot, R. Geesink, C.P.A.T. Klein, P. Serekian, J. Biomed. Mater. Res. 21 (1987) 1375. [4] J. Weng, Q. Liu, J.G.C. Wolke, X. Zhang, K. de Groot, Biomaterials 18 (1997) 1027. [5] J.H. Chern Lin, M.L. Liu, C.P. Ju, J. Mater. Sci.: Mater. Med. 28 (1994) 1. [6] J.G.C. Wolke, J.P.C.M. van der Waerden, K. de Groot, J.A. Jansen, Biomaterials 18 (1997) 483. [7] S.J. Ding, C.P. Ju, J.H. Chern Lin, J. Biomed. Mater. Res. 44 (1999) 266. [8] C.K. Wang, J.H. Chern Lin, C.P. Ju, H.C. Ong, R.P.H. Chang, Biomaterials 18 (1997) 1331. [9] C.M. Lopatin, V. Pizziconi, T.L. Alford, T. Laursen, Thin Solid Films 326 (1998) 227. [10] P. Ducheyne, S. Radin, M. Heughebaert, J.C. Heughebaert, Biomaterials 11 (1990) 244. [11] M. Gottlander, C.B. Hohansson, A. Wennerberg, T. Albrektsson, S. Ducheyne, P. Ducheyne, Biomaterials 18 (1997) 551. [12] M. Shirkhanzadeh, J. Mater. Sci. Lett. 10 (1991) 1415. [13] T.V. Vijayaraghavan, A. Bensalem, J. Mater. Sci. Lett. 13 (1994) 1782. [14] S. Ban, S. Maruno, Biomaterials 16 (1995) 977. [15] M. Shirkhanzadeh, J. Mater. Sci.: Mater. Med. 9 (1998) 67. [16] M. Shirkhanzadeh, M. Azadegan, Mater. Lett. 15 (1993) 392. [17] J. Huaxia, C.B. Ponton, P.M. Marquis, J. Mater. Sci.: Mater. Med. 3 (1992) 283. [18] S.R. Radin, P. Ducheyne, J. Mater. Sci.: Mater. Med. 3 (1992) 33. [19] J. Redepenning, J.P. McIsaac, Chem. Mater. 2 (1990) 625. [20] R.Z. LeGeros, Prog. Cryst. Growth Charact. 4 (1981) 1. [21] A.S. Posner, Clin. Orthop. Relat. Res. 200 (1985) 87.
160
M.C. Kuo, S.K. Yen / Materials Science and Engineering C 20 (2002) 153–160
[22] L.M. Rodrigues-Lorenzo, M. Vallet-Regi, J.M.F. Ferreira, Biomaterials 22 (2001) 583. [23] G. Daculsi, Biomaterials 19 (1998) 1473. [24] N. Verdonschot, M. Dalstra, R. Huskes, in: G. Bergmann, F. Graidren, A. Roklmann (Eds.), Implantable Telemetry in Orthopedics, Berlin Free University, Berlin, Germany, 1990, p. 249.
[25] J.H. Chern Lin, H.J. Lin, S.J. Ding, C.P. Ju, Mater. Chem. Phys. 64 (2000) 229. [26] L.-D. Piveteau, B. Gasser, L. Schlapbach, Biomaterials 21 (2000) 2193.