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Magnesium substituted hydroxyapatite formation on (Ti,Mg)N coatings produced by cathodic arc PVD technique
Materials Science and Engineering C 33 (2013) 4337–4342
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Magnesium substituted hydroxyapatite formation on (Ti,Mg)N coatings produced by cathodic arc PVD technique Sakip Onder a,b, Fatma Nese Kok b,c, Kursat Kazmanli a,⁎, Mustafa Urgen a a b c
Department of Metallurgical and Materials Engineering, Istanbul Technical University, 34469, Maslak, Istanbul, Turkey Molecular Biology-Genetics and Biotechnology Program (MOBGAM), Istanbul Technical University, 34469, Maslak, Istanbul, Turkey Department of Molecular Biology and Genetics, Istanbul Technical University, 34469, Maslak, Istanbul, Turkey
a r t i c l e
i n f o
Article history: Received 26 March 2013 Received in revised form 21 May 2013 Accepted 19 June 2013 Available online 28 June 2013 Keywords: Hydroxyapatite Magnesium (Ti,Mg)N coating TiN coating Cathodic Arc PVD Technique
a b s t r a c t In this study, formation of magnesium substituted hydroxyapatite (Ca10−xMgx(PO4)6(OH)2) on (Ti,Mg)N and TiN coating surfaces were investigated. The (Ti1−x,Mgx)N (x = 0.064) coatings were deposited on titanium substrates by using cathodic arc physical vapor deposition technique. TiN coated grade 2 titanium substrates were used as reference to understand the role of magnesium on hydroxyapatite (HA) formation. The HA formation experiments was carried out in simulated body fluids (SBF) with three different concentrations (1X SBF, 5X SBF and 5X SBF without magnesium ions) at 37 °C. The coatings and hydroxyapatite films formed were characterized by scanning electron microscope (SEM), X-ray diffraction (XRD) and FTIR Spectroscopy techniques. The energy dispersive X-ray spectroscopy (EDS) analyses and XRD investigations of the coatings indicated that magnesium was incorporated in the TiN structure rather than forming a separate phase. The comparison between the TiN and (Ti, Mg)N coatings showed that the presence of magnesium in TiN structure facilitated magnesium substituted HA formation on the surface. The (Ti,Mg)N coatings can potentially be used to accelerate the HA formation in vivo conditions without any prior hydroxyapatite coating procedure. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Hydroxyapatite (Ca10(PO4)6(OH)2) is a well-known biocompatible material that has the most similar chemical structure to that of bone minerals [1,2]. This makes hydroxyapatite (HA) an attractive biomaterial in hard-tissue applications. HA could improve the integration of the implant to the surrounding tissue by providing strong bonds between implant material and bone tissue interface and, thus, accelerating bone growth from the surface [3]. HA coated titanium surfaces exhibit better osteoblast cell attachment and proliferation compared to the unmodified surfaces [4–12]. Moreover, it has been reported that doping HA structures with different ions such as Mg2+ could improve mineralization of calcified tissues and indirectly influence mineral metabolism [13–17]. Magnesium is present in the structure of naturally formed HA and forms part of the HA structure by partially substituting for calcium. In human body, Mg2+ ion is the most abundant cation [18] and cofactor for many enzymes some of which are known to induce the proliferation of osteoblast cells [19]. Magnesium, therefore, has an active role on “new bone tissue generation” process [20,21]. In the literature, HA has been synthesized from aqueous solutions in powder form by different techniques such as sol-gel [16] and
⁎ Corresponding author. Tel.: +90 212 285 3537; fax: +90 212 285 3427. E-mail address: [email protected] (K. Kazmanli). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.06.027
mechanochemical/mechanochemical–hydrothermal syntheses [22–29]. After the synthesis process, a subsequent heat treatment was conducted to complete the HA formation reactions and crystallization. The heat treatment process conducted in a temperature range of 300–1000 °C may lead to the formation of undesired whitlockite phase [30]. For HA deposition on implant surfaces, various coating methods such as pulsed laser ablation, sputter deposition, electrophoretic deposition, thermal and plasma spray techniques are used [31–52]. In these deposition processes, either HA powders or targets produced by chemical or mechanochemical methods are used as source materials. Sol-gel method is another technique used for the deposition of HA. The structure and chemistry of HA films show a strong dependence on the deposition technique used. Amorphous structure or low crystallinity of deposited HA films remains to be the main issue on the production of HA coatings since it directly affects the biological performance. For studying the properties and kinetics of HA formation on implant surfaces, simulated body fluids (SBF) are widely used. HA can naturally be formed on implant surfaces in these solutions simulating the physiological environment. The SBF experiments are extensively utilized for the evaluation of nucleation kinetics and properties of HA formed on different implant surfaces. Detailed reviews on this subject can be found in the literature [18,21,53]. Since SBF contain Mg ions, the substitution for Ca ions is expected to some extent during the HA formation process. In this context, Suchanek et al. [21] and Barrere et al. [54] synthesized Mg-HA with using excessive amount of Mg and studied consequent structural and chemical changes on HA. Although this has given
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a)
Fig. 2. (a) Full range XRD patterns of TiN and (Ti,Mg)N coatings, (b) Detailed view of XRD patterns around diffraction peak of (111) crystal plane.
Fig 1. Schematic of the coating system.
valuable insights on the effect of Mg ion concentration on HA formation and properties, high magnesium containing solutions fail to simulate the actual body environments so could not be used to comment on HA formation in vivo conditions. Relying on the positive role of Mg2+ ions on osteo-integration of implants and growth kinetics of HA, two approaches can be used for the increased availability of Mg2+ ions on implant surfaces: 1. Using Mg containing alloys as implant materials that have ability to release Mg in a controlled manner into the body fluid. However, pure magnesium and Mg based alloys are not suitable materials for implants since they have low corrosion resistance in most electrolyte mediums [18,20] limiting their usage in the human body environment. 2. Deposition of bio–compatible coatings containing a desired amount of magnesium on implant materials to give the implant surface a “magnesium releasing ability” in a controlled manner. The methodology and hypothesis of this study is based on the second approach; namely modifying the implant surfaces with magnesium containing coating that will contribute to HA formation on the implant surfaces in the body environment. For this purpose, a biocompatible coating that can be alloyed with magnesium is required. As the biocompatible coating, TiN, which is a widely used coating in biomedical applications, was selected. In order to dope TiN with magnesium, cathodic arc physical vapor deposition technique was utilized. In the literature there are few studies [55,56] on magnesium containing TiN coatings and these studies were mainly focused on optical [55] and high temperature oxidation behaviors [56]. These studies showed that magnesium atoms substituted the titanium atoms in
Table 1 Chemical Composition of 1X and 5X simulated body fluids. 1X SBF
b)
5X SBF
Order
Reagent
Amount (gr) in 40 ml solution
Reagent
Amount (gr) in 40 ml solution
#1 #2 #3 #4 #5 #6 #7 #8
NaCl NaHCO3 KCl K2HPO4 .3H2O MgCl2.6H2O HCl (1 M) CaCl2 Na2SO4
0.319 0.013 0.009 0.009 0.01 1.48 ml 0.01 0.003
CaCl2 MgCl2.6H2O NaHCO3 K2HPO4 .3H2O Na2SO4 KCl NaCl HCl (1 M)
0.056 0.06 0.07 0.045 0.015 0.045 1.6 Used for adjusting pH to 6
#9
(CH2OH)3CNH2
0.24
the cubic TiN structure forming (Ti1−x,Mgx)N phase even at high magnesium concentrations (x = 0.67). In the present study, the role of magnesium released from (Ti,Mg) N coatings on nucleation and growth of HA in SBF was investigated. In order to compare and understand the effect of magnesium on HA formation, magnesium free TiN coatings were also deposited and tested in two different SBF solutions with or without magnesium ions. The surface of the coatings and the HA films formed in SBF were characterized by SEM, XRD and FTIR techniques. 2. Material and methods 2.1. Deposition and characterization of (Ti,Mg)N and TiN coatings The coatings were deposited on grade 2 Ti plates (15x15x1 mm) by using cathodic arc physical vapor deposition technique. Our previous results [57] on the corrosion resistance of (Ti1–x,Mgx)N revealed that coatings with more than x = 0.12 Mg were not stable in chloride containing environments and better corrosion resistance has been achieved with x = 0.057. The parameters of the deposition process were, thus, selected to produce coatings containing x ≈ 0.06 Mg. Before the deposition, the substrates were polished using 1000, 2500 and 4000-grit SiC papers sequentially. Shortly after the polishing, the substrates were ultrasonically cleaned in acetone and ethanol baths for 10 minutes. After cleaning, the substrates were placed in the vacuum chamber then the chamber was evacuated down to 5 × 10−3 Pa. Since evaporation rate of Mg is higher than that of Ti, two Ti and one Mg cathodes were used to adjust Mg content of the coatings. The arc currents were 75 A for each Ti cathodes and 40 A for the Mg cathode. Before the coating process, the substrate surfaces were heated and etched by Ti ions produced as a result of cathodic arc evaporation. In the ion heating/etching process, a series of high bias voltages were applied on the substrates (–600, –800 and –1000 V in 30 sec for each step). The deposition process was conducted at 1 Pa N2 pressure and at –250 V bias voltages. In order to produce homogeneous coatings, the samples were rotated so that they could be exposed to each cathode for equal time intervals. Total coating time was 1 hour for each coating. The schematic of the coating system was given in Fig. 1. In order to understand the role of Mg on HA formation, a magnesium free TiN coating was also deposited. The magnesium content of the coatings was analyzed by using scanning electron microscope (SEM) with energy dispersive X-ray spectrometer (EDS) (JEOL JSM 7000F). The phase structure of the coatings and the HA deposits were analyzed by using glancing angle X-ray diffractometer with thin film attachment (Philips Model
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Fig. 3. SEM micrographs of HA deposits on TiN and (Ti,Mg)N after 4 weeks in 1X SBF.
PW3710). The Cu–Kα radiation and “θ scan method” with fixed incidence angle of 0.5° were used. After SBF experiments, the deposits formed on the coatings were investigated by using FTIR spectrometer and SEM. 2.2. Formation of hydroxyapatite in simulated body fluid Two different simulated body fluid concentrations (1X SBF and 5X SBF) were prepared as described in the references [58] and [59] (Table 1). The final pH values of 1X SBF and 5X SBF were adjusted to 7.4 and 6.5, respectively at 37 °C by using 1 M hydrochloric acid
or 1 M sodium hydroxide. The samples were incubated in 1X SBF for 1, 2, 3 and 4 weeks in order to monitor the HA formation. In experiments conducted in 5X SBF, HA film formation was accelerated so the samples were kept in the solution for only 3 days. TiN and (Ti,Mg)N coated samples were also tested for 3 days in magnesium free 5X SBF solution to investigate the role of Mg on the kinetics of HA formation. The HA deposits formed on TiN and (Ti,Mg)N coated samples in 1X, 5X and Mg free 5X SBF solutions were characterized by using FTIR spectroscopy and XRD techniques. 3. Results and discussions 3.1. Characterization of (Ti,Mg)N and TiN coatings
TiN peak (111)
Before SBF
TiN
(Ti,Mg)N Counts (a.u)
After SBF
The thickness of the coatings as determined from their fracture cross sections were 1.1 μm and 1.2 μm for TiN and (Ti,Mg)N, respectively. The EDS results showed that the (Ti1−x,Mgx)N coating contained Mg equivalent to x = 0.064. In the X-ray diffraction patterns of both coatings, characteristic TiN peaks at 36.37° (111), 42.43 (200), 61.63° (220) and 73.85° (311) were observed as shown in Fig. 2a. Furthermore, the peaks corresponding to magnesium or magnesium nitride phases were not present in the XRD pattern of the (Ti,Mg)N sample. Lattice parameters of the TiN and (Ti,Mg)N coatings were calculated from the XRD patterns by using “Philips X'PERT Plus Crystallography and Rietveld Analysis Software”. The calculated lattice parameters of the TiN and (Ti,Mg)N coatings were a = 4.248 Å and a = 4.253 Å, respectively. Magnesium addition shifted the TiN peaks to lower diffraction angles (Fig. 2b) expanding
1 week
2 weeks
3 weeks 4 weeks 35
36
37
38
2 Theta(deg) Fig. 4. XRD patterns of (Ti,Mg)N coatings after 1X SBF test.
Fig. 5. FTIR spectra of deposits formed on TiN and (Ti,Mg)N coatings in 1X SBF.
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Fig. 6. SEM micrographs of (a) (Ti,Mg)N and (b) TiN coatings after 5X SBF tests.
In 1X SBF, HA formation on the (Ti,Mg)N coating was initiated after the first week of immersion. HA deposits on the TiN coating, on the other hand, reached to an observable size only after 4th week. Higher amount of HA deposits on the (Ti,Mg)N could clearly been observed at the end of 4th week in the SEM micrograph (Fig. 3). The HA deposits, however, did not seem to homogeneously cover the whole surface after 4 weeks and the HA derived peaks were not observed in the XRD patterns of both coatings. Nevertheless, the strongest (111) TiN peak of the (Ti,Mg)N coating shifted to the diffraction angle of magnesium free TiN (111) peak after 1st week of immersion (Fig. 4). The shift of TiN peaks could be associated with the dissolution of Mg from the (Ti,Mg)N structure. After the first week, the FTIR spectrum (Fig. 5) of the (Ti,Mg)N sample exhibited PO−3 peaks around 553 and 4 1021 cm−1. Although PO−3 peaks in the spectrum of TiN started to 4 evolve after second week, their intensity were not comparable to those of the (Ti,Mg)N even after 4 weeks. Since the HA deposits formed in 1X SBF did not continuously cover the surfaces after 4 weeks, the HA formation experiments were
repeated in 5X SBF solution for a duration of 3 days. The SEM results showed that incubation of the samples in this solution led to a complete coverage of both TiN and (Ti,Mg)N surfaces with HA deposits after 3 days (Fig. 6). It could be clearly seen that the HA nodules on the (Ti,Mg)N surface developed to bigger sizes compared to those on the TiN surface (i.e. 2–3 times larger). The XRD patterns of the HA deposits on both coatings exhibited characteristic HA peaks (Fig. 7). FTIR spectra of the HA deposits on both coatings were given in Fig. 8. In the spectra, PO−3 group peaks at 475, 561, 600, 958, 1014 and 4 1115 cm−1, HPO−2 group peak at 867 cm−1, CO−2 peak at 1571 cm−1 4 3 −1 and H2O peak at 1634 cm are characteristic peaks of a hydroxyapatite structure. These results were in accordance with the literature; Suchanek et al. [21] and Hanifi et al. [60] reported that PO−3 peaks be4 came smaller with increasing amount of Mg in HA. Suchanek also reported that HPO−2 peak at 870 cm−1 evolved with higher magne4 sium contents in the HA [21]. In our study, PO−3 peak at 475 cm−1 4 −1 disappeared and CO−2 peaks at 1458 and 1571 cm became more 3 pronounced in the spectrum of the (Ti,Mg)N sample (Fig. 8a) when compared with the corresponding peaks of the TiN sample (Fig. 8b). Hanifi et al., and Suchanek et al. also reported that substitution of CO−2 3 for PO−3 was particularly observed in the Mg substituted HA structure. 4 −3 The similarities of CO−2 and HPO−2 peaks with the existing 3 , PO4 4 literature [21,60] is a strong indication of the Mg substituted HA formation on the (Ti,Mg)N coated sample. It could therefore be stated that the
Fig. 7. XRD analysis of TiN and (Ti,Mg)N coatings after 5X SBF tests.
Fig. 8. FTIR spectra of (a) (Ti,Mg)N and (b) TiN coatings after 5X SBF tests.
the TiN lattice by 0.17 %. Hodroj et al. [56] and Fenker et al. [55] reported similar results showing that magnesium substitution for titanium atom in TiN structure resulted in lattice expansion of TiN phase because of higher atomic radius of magnesium. 3.2. Hydroxyapatite formation in Simulated Body Fluids
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Fig. 9. SEM micrographs of (a) (Ti,Mg)N and (b) TiN coatings after Mg-free 5X SBF tests.
incorporation of Mg into the TiN structure promoted the Mg substituted hydroxyapatite formation on the coating. For revealing the role of Mg2+ ions on the growth of HA for both coatings, the 5X SBF tests were repeated in the SBF without Mg2+ ions. After 3 days of incubation, there was no HA formation on the TiN sample (Fig. 9b). Barrere et al. [54] showed the key role of Mg2+ ion on the formation of HA on Ti6Al4V substrates. They immersed the Ti6Al4V substrates in 5X SBF with and without Mg2+ ions. After 24 hours of immersion, they observed that HA deposits could only be formed in the solution containing Mg2+ ions. The results of our experiments further verified the importance of magnesium presence in solution during HA formation, this time in TiN coated samples. It should, however, be noted that HA formation was observed on the surface of (Ti,Mg)N coating in 5X SBF solution even in the absence of Mg2+ ions (SEM micrographs, Fig 9a). The HA derived peaks in the FTIR spectrum (Fig. 10) of the (Ti,Mg)N sample also showed the HA formation on the (Ti,Mg)N coating in this solution. These results clearly showed that the Mg in the (Ti,Mg)N coating promoted HA formation by releasing magnesium ions to the “SBF/ coating” interface.
cathodic arc PVD technique. The XRD patterns of the (Ti1–x,Mgx)N coatings indicated that addition of Mg equivalent to x = 0.064 did not form a separate Mg or magnesium nitride phases, instead, Mg incorporated in the TiN structure shifting the TiN diffraction peaks in the XRD pattern. In consequence of Mg dissolution into the simulated body fluid during the HA formation process, Mg-substituted HA formation was facilitated. In other words, the Mg doped TiN structure acted as a magnesium source promoting the Mg-substituted HA formation in simulated body fluids. This encouraging result offers the potential to produce hard tissue implants with the ability to induce on-site HA production on the surface, allowing improved tissue integration.
4. Conclusion
References
The effect of Mg on HA formation on TiN and (Ti,Mg)N coatings was studied. The coatings were deposited on Ti substrates by using
[1] B. Leukers, H. Gülkan, S.H. Irsen, S. Milz, C. Tille, M. Schieker, H. Seitz, Hydroxyapatite scaffolds for bone tissue engineering made by 3D printing, J. Mater. Sci. Mater. Med. 16 (2005) 1121–1124. [2] S. Teixeira, M.A. Rodriguez, P. Pena, A.H. De Aza, S. De Aza, M.P. Ferraz, F.J. Monteiro, Physical characterization of hydroxyapatite porous scaffolds for tissue engineering, Mater. Sci. Eng. C 29 (2009) 1510–1514. [3] P.K. Bajpai, C.M. Fuchs, Development of a hydroxyapatite bone grout, in: C.W. (Ed.), Proceedings of the First Annual Scientific Session of the Academy of Surgical Research, Pergamon Press, NY, 1985. [4] T.J. Webster, R.W. Siegel, R. Bizios, Bioceramics, World Scientific, New York, 1998. [5] T.J. Webster, R.W. Siegel, R. Bizios, Osteoblast adhesion on nanophase ceramics, Biomaterials 20 (1999) 1222–1227. [6] Y. Liu, T. Jiang, Y. Zhou, Z. Zhang, Z. Wang, H. Tong, X. Shen, Y. Wang, Evaluation of the attachment, proliferation, and differentiation of osteoblast on a calcium carbonate coating on titanium surface, Mater. Sci. Eng. C 31 (2011) 1055–1061. [7] W.K. Lee, S.M. Lee, H.M. Kim, Effect of surface morphology of calcium phosphate on osteoblast-like HOS cell responses, J. Ind. Eng. Chem. 15 (2009) 5. [8] J. Huang, S.M. Best, W. Bonfield, T. Buckland, Development and characterization of titanium- containing hydroxyapatite for medical applications, Acta Biomater. (2010) 241–249. [9] N. Sato, K. Kubo, M. Yamada, N. Hori, T. Suzuki, H. Maeda, T. Ogawa, Osteoblast mechanoresponses on Ti with different surface topographies, J. Dent. Res. 88 (2009) 812–816. [10] I. Wall, N. Donos, K. Carlqvist, F. Jones, P. Brett, Modified titanium surfaces promote accelerated osteogenic differentiation of mesenchymal stromal cells in vitro, Bone 45 (2009) 17–26. [11] K.Y. Cai, Y. Hu, K.D. Jandt, Y.L. Wang, Surface modification of titanium thin film with chitosan via electrostatic self-assembly technique and its influence on osteoblast growth behavior, J. Mater. Sci. Mater. Med. 19 (2008) 499–506. [12] M.D. Ball, S. Downes, C.A. Scotchford, E.N. Antonov, V.N. Bagratashvili, V.K. Popov, W.J. Lo, D.M. Grant, S.M. Howdle, Osteoblast growth on titanium foils coated with hydroxyapatite by pulsed laser ablation, Biomaterials 22 (2001) 337–347. [13] F. Ren, et al., Synthesis, characterization and ab initio simulation of magnesium substituted hydroxyapatite, Acta Biomater. 6 (2010) 2787–2796.
Fig. 10. FTIR spectra of (a) TiN and (b) (Ti,Mg)N coatings after Mg-free 5X SBF tests.
Acknowledgments This study was supported by The Scientific and Research Council of Turkey (TUBITAK –Project #112M339). The authors gratefully thank Mehmet Dokur for his valuable helps in FTIR analyses and ITU Graduate School of Science, Engineering and Technology for ITU-BAP Project #34525.
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S. Onder et al. / Materials Science and Engineering C 33 (2013) 4337–4342
[14] B. Bracci, et al., Effect Of Mg(2+), Sr(2+), And Mn(2+) on the chemico–physical and in vitro biological properties of calcium phosphate biomimetic coatings, J. Inorg. Biochem. 103 (2009) 1666–1674. [15] S.D. Garcia, D. Laurencin, J.M. Krafft, S. Casale, M.E. Smith, H.L. Pernot, G. Costentin, Influence of magnesium substitution on the basic properties of hydroxyapatites, J. Phys. Chem. C 115 (2011) 24317–24327. [16] A. Gozalian, A. Behnamghader, M. Daliri, A. Moshkforoush, Synthesis and thermal behavior of Mg-doped calcium phosphate nanopowders via the sol gel method, Sci. Iran 6 (2011) 1614–1622. [17] I. Cacciotti, et al., Mg-substituted hydroxyapatite nanopowders: synthesis, thermal stability and sintering behavior, J. Eur. Ceram. Soc. 14 (2009) 2969–2978. [18] L. Li, J. Gao, Y. Wang, Evaluation of cyto-toxicity and corrosion behavior of alkali heat-treated magnesium in simulated body fluid, Surf. Coat. Technol. 185 (2004) 92–98. [19] J. Takaya, H. Higashino, Y. Kobayashi, Intracellular magnesium and insulin resistance, Magnes. Res. 17 (2004) 126–136. [20] M.P. Staiger, A.M. Pietak, J. Huadmai, G. Dias, Magnesium and its alloys as orthopedic biomaterials: a review, Biomaterials 27 (2006) 1728–1734. [21] W.L. Suchanek, K. Byrappa, P. Shuk, R.E. Riman, V.F. Janas, K.S. TenHuisen, Preparation of magnesium-substituted hydroxyapatite powders by the mechanochemical– hydrothermal method, Biomaterials 25 (2004) 4647–4657. [22] M. Toriyama, S. Kawamura, Synthesis of b-tricalcium phosphate by use of wet milling, J. Ceram. Soc. Jpn. 94 (1986) 1004–1007. [23] M. Toriyama, S. Kawamura, Sinterable powder of mechanochemically synthetic b-tricalcium phosphate, J. Ceram. Soc. Jpn. 95 (1987) 741–745. [24] M. Toriyama, S. Kawamura, Y. Ito, H. Nagae, Thermal change of calcium deficient apatite obtained by mechanochemical treatment, J. Ceram. Soc. Jpn. 97 (1989) 554–558. [25] M. Otsuka, Y. Matsuda, J. Hsu, J.L. Fox, W.I. Higuchi, Mechanochemical synthesis of bioactive material: effect of environmental conditions on the phase transformation of calcium phosphates during grinding, Biomed. Mater. Eng. 4 (1994) 357–362. [26] M. Toriyama, A. Ravaglioli, A. Krajewski, C. Galassi, E. Roncari, A. Piancastelli, Slip casting of mechanochemically synthesized hydroxyapatite, J. Mater. Sci. 30 (1995) 3216–3221. [27] M. Toriyama, A. Ravaglioli, A. Krajewski, G. Celotti, A. Piancastelli, Synthesis of hydroxyapatite-based powders by mechanochemical method and their sintering, J. Eur. Ceram. Soc. 16 (1996) 429–436. [28] Y. Yokogawa, M. Toriyama, Y. Kawamoto, T. Suzuki, K. Nishizawa, F. Nagata, M.R. Mucalo, Preparation of calcium strontium apatite through mechanochemical method, Chem. Lett. 1 (1996) 91–92. [29] J. Liao, K. Ono, G. Kanayama, T. Isobe, M. Senna, Preparation of tricalcium phosphate by applying grinding technique and choosing hydroxy-containing initial materials, Chem. Sustainable Dev. 6 (1998) 233–236. [30] I.V. Fadeev, et al., Synthesis and structure of magnesium-substituted hydroxyapatite, Inorg. Mater. 39 (2003) 947–950. [31] Y. Zhao, C. Chen, D. Wang, Surf. Rev. Lett. 12 (2005) 401–408. [32] E. Van der Wal, A.M. Vredenberg, P.J. Ter Brugge, J.G.C. Wolke, J.A. Jansen, Biomaterials 27 (2006) 1333–1340. [33] M. Inagaki, Y. Yokogawa, T. Kameyama, Surf. Coat. Technol. 173 (2003) 1–8. [34] M. Roy, A. Bandyopadhyay, S. Bose, Induction plasma sprayed Sr and Mg doped nano hydroxyapatite coatings on Ti for bone implant, J. Biomed. Mater. Res. B 99B (2011) 2. [35] T. Huang, et al., Nanostructured Si, Mg, CO3 (2-) substituted hydroxyapatite coatings deposited by liquid precursor plasma spraying: synthesis and characterization, J. Therm. Spray Technol. 4 (2011) 829–836. [36] Y. Yunzhi, L.O. Kyo-Han Kim Joo, A review on calcium phosphate coatings produced using a sputtering process—an alternative to plasma spraying, Biomaterials 26 (2005) 327–337. [37] K.A. Gross, C.C. Berndt, Thermal processing of hydroxyapatite for coating production, J. Biomed. Mater. Res. 39 (1998) 580–587.
[38] K.A. Gross, C.C. Berndt, H. Herman, Amorphous phase formation in plasma-sprayed hydroxyapatite coatings, J. Biomed. Mater. Res. 39 (1998) 407–414. [39] H. Li, K.A. Khor, P. Cheang, Titanium dioxide reinforced hydroxyapatite coatings deposited by high velocity oxy-fuel (HVOF) spray, Biomaterials 23 (2002) 85–91. [40] Y. Yang, J.L. Ong, Bond strength, compositional and structural properties of HA coating on Ti substrate, ZrO2-coated Ti substrate, and TPS-coated Ti substrate, J. Biomed. Mater. Res. 64 (2003) 509–516. [41] S. Tao, H. Ji, C. Ding, Effect of vapor–flame treatment on plasma sprayed hydroxyapatite coatings, J. Biomed. Mater. Res. 52 (2000) 572–575. [42] X. Weichang, L. Xuanyong, Z. Xuebin, D. Chuanxian, Effect of hydroxyapatite coating crystallinity on dissolution and osseointegration in vivo, J. Biomed. Mater. Res. A 74A (2005) 553–561. [43] S. Ding, C. Ju, J.C. Lin, Characterization of hydroxyapatite and titanium coatings sputtered on Ti-6Al-4 V substrate, J. Biomed. Mater. Res. 44 (1999) 266–279. [44] C. Massaro, M.A. Baker, F. Cosentino, P.A. Ramires, S. Klose, E. Milella, Surface and biological evaluation of hydroxyapatite-based coatings on titanium deposited by different techniques, J. Biomed. Mater. Res. (Appl. Biomater.) 58 (2001) 651–657. [45] V. Nelea, C. Morosanu, M. Iliescu, I.N. Mihailescu, Hydroxyapatite thin films grown by pulsed laser deposition and radio-frequency magnetron sputtering: comparative study, Appl. Surf. Sci. 228 (2004) 346–356. [46] H. Zeng, W.R. Lacefield, The study of surface transformation of pulsed laser deposited hydroxyapatite coatings, J. Biomed. Mater. Res. 50 (2000) 239–247. [47] G. Jiang, D. Shi, Coating of hydroxyapatite on highly porous Al2O3 substrate for bone substitutes, J. Biomed. Mater. Res. (Appl. Biomater.) 43 (1998) 77–81. [48] Y. Han, K. Xu, J. Lu, Z. Wu, The structural characteristics and mechanical behaviors of nonstoichiometric apatite coatings sintered in air atmosphere, J. Biomed. Mater. Res. 45 (1999) 198–203. [49] Y. Han, T. Fu, J. Lu, K. Xu, Characterization and stability of hydroxyapatite coatings prepared by an electrodeposition and alkaline-treatment process, J. Biomed. Mater. Res. 54 (2001) 96–101. [50] L.A. De Sena, M.C. de Andrade, A.M. Rossi, G.D.A. Soares, Hydroxyapatite deposition by electrophoresis on titanium sheets with different surface finishing, J. Biomed. Mater. Res. 60 (2002) 1–7. [51] P. Habibovic, F. Barrere, C.A. van Bliterswijk, K. de Groot, P. Layrolle, Biomimetic hydroxyapatite coating on metal implants, J. Am. Ceram. Soc. 85 (2002) 517–522. [52] A.L. Oliveira, C. Elvira, R.L. Reis, B. Vazquez, J. San Roman, Surface modification tailors the characteristics of biomimetic nucleated on starch-based polymers, J. Mater. Sci. Mater. Med. 10 (1995) 827–835. [53] C. Ohtsuki, T. Kokubo, T. Yamamuro, Mechanism of HA formation of CaO-SiO2-P2O5 glasses in simulated body fluid, J. Non-Cryst. Solids 43 (1992) 84–92. [54] F. Barrere, et al., Nucleation of biomimetic Ca-P coatings on Ti6Al4V from a SBF x 5 solution: influence of magnesium, Biomaterials 23 (2002) 2211–2220. [55] M. Fenker, M. Balzer, H. Kappl, O. Banakh, Some properties of (Ti, Mg)N thin films deposited by reactive dc magnetron sputtering, Surf. Coat. Technol. 200 (2005) 227–231. [56] A. Hodroj, O. Chaix-Pluchery, P. Steyer, J.F. Pierson, Oxidation resistance of decorative (Ti,Mg)N coatings deposited by hybrid cathodic arc evaporation-magnetron sputtering process, Surf. Coat. Technol. (2011) 4547–4553. [57] B.S. Demirel, Iyon Demeti Destekli Manyetik Alanda Sıçratma Yöntemiyle (Ti,Mg) N İnce Filmlerin Üretilmesi, Yapısal ve Biyoaktivite Özelliklerinin Belirlenmesi, (MSc Thesis) Istanbul Technical University, Graduate School of Science, Engineering and Technology, Istanbul, 2003. [58] T. Kokubo, Apatite formation on surfaces o f ceramics, metals and polymers in body environment, Acta Biomater. 146 (1998) 2519–2527. [59] Y.F. Chou, et al., The effect of pH on the structural evolution of accelerated biomimetic apatite, Biomaterials 25 (2004) 5323–5331. [60] A. Hanifi, et al., Mg2 + substituted calcium phosphate nano particles synthesis for non viral gene delivery application, J. Mater. Sci. Mater. Med. 21 (2010) 2393–2401.
Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Magnesium substituted hydroxyapatite formation on (Ti,Mg)N coatings produced by cathodic arc PVD technique Sakip Onder a,b, Fatma Nese Kok b,c, Kursat Kazmanli a,⁎, Mustafa Urgen a a b c
Department of Metallurgical and Materials Engineering, Istanbul Technical University, 34469, Maslak, Istanbul, Turkey Molecular Biology-Genetics and Biotechnology Program (MOBGAM), Istanbul Technical University, 34469, Maslak, Istanbul, Turkey Department of Molecular Biology and Genetics, Istanbul Technical University, 34469, Maslak, Istanbul, Turkey
a r t i c l e
i n f o
Article history: Received 26 March 2013 Received in revised form 21 May 2013 Accepted 19 June 2013 Available online 28 June 2013 Keywords: Hydroxyapatite Magnesium (Ti,Mg)N coating TiN coating Cathodic Arc PVD Technique
a b s t r a c t In this study, formation of magnesium substituted hydroxyapatite (Ca10−xMgx(PO4)6(OH)2) on (Ti,Mg)N and TiN coating surfaces were investigated. The (Ti1−x,Mgx)N (x = 0.064) coatings were deposited on titanium substrates by using cathodic arc physical vapor deposition technique. TiN coated grade 2 titanium substrates were used as reference to understand the role of magnesium on hydroxyapatite (HA) formation. The HA formation experiments was carried out in simulated body fluids (SBF) with three different concentrations (1X SBF, 5X SBF and 5X SBF without magnesium ions) at 37 °C. The coatings and hydroxyapatite films formed were characterized by scanning electron microscope (SEM), X-ray diffraction (XRD) and FTIR Spectroscopy techniques. The energy dispersive X-ray spectroscopy (EDS) analyses and XRD investigations of the coatings indicated that magnesium was incorporated in the TiN structure rather than forming a separate phase. The comparison between the TiN and (Ti, Mg)N coatings showed that the presence of magnesium in TiN structure facilitated magnesium substituted HA formation on the surface. The (Ti,Mg)N coatings can potentially be used to accelerate the HA formation in vivo conditions without any prior hydroxyapatite coating procedure. © 2013 Elsevier B.V. All rights reserved.
1. Introduction Hydroxyapatite (Ca10(PO4)6(OH)2) is a well-known biocompatible material that has the most similar chemical structure to that of bone minerals [1,2]. This makes hydroxyapatite (HA) an attractive biomaterial in hard-tissue applications. HA could improve the integration of the implant to the surrounding tissue by providing strong bonds between implant material and bone tissue interface and, thus, accelerating bone growth from the surface [3]. HA coated titanium surfaces exhibit better osteoblast cell attachment and proliferation compared to the unmodified surfaces [4–12]. Moreover, it has been reported that doping HA structures with different ions such as Mg2+ could improve mineralization of calcified tissues and indirectly influence mineral metabolism [13–17]. Magnesium is present in the structure of naturally formed HA and forms part of the HA structure by partially substituting for calcium. In human body, Mg2+ ion is the most abundant cation [18] and cofactor for many enzymes some of which are known to induce the proliferation of osteoblast cells [19]. Magnesium, therefore, has an active role on “new bone tissue generation” process [20,21]. In the literature, HA has been synthesized from aqueous solutions in powder form by different techniques such as sol-gel [16] and
⁎ Corresponding author. Tel.: +90 212 285 3537; fax: +90 212 285 3427. E-mail address: [email protected] (K. Kazmanli). 0928-4931/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.msec.2013.06.027
mechanochemical/mechanochemical–hydrothermal syntheses [22–29]. After the synthesis process, a subsequent heat treatment was conducted to complete the HA formation reactions and crystallization. The heat treatment process conducted in a temperature range of 300–1000 °C may lead to the formation of undesired whitlockite phase [30]. For HA deposition on implant surfaces, various coating methods such as pulsed laser ablation, sputter deposition, electrophoretic deposition, thermal and plasma spray techniques are used [31–52]. In these deposition processes, either HA powders or targets produced by chemical or mechanochemical methods are used as source materials. Sol-gel method is another technique used for the deposition of HA. The structure and chemistry of HA films show a strong dependence on the deposition technique used. Amorphous structure or low crystallinity of deposited HA films remains to be the main issue on the production of HA coatings since it directly affects the biological performance. For studying the properties and kinetics of HA formation on implant surfaces, simulated body fluids (SBF) are widely used. HA can naturally be formed on implant surfaces in these solutions simulating the physiological environment. The SBF experiments are extensively utilized for the evaluation of nucleation kinetics and properties of HA formed on different implant surfaces. Detailed reviews on this subject can be found in the literature [18,21,53]. Since SBF contain Mg ions, the substitution for Ca ions is expected to some extent during the HA formation process. In this context, Suchanek et al. [21] and Barrere et al. [54] synthesized Mg-HA with using excessive amount of Mg and studied consequent structural and chemical changes on HA. Although this has given
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a)
Fig. 2. (a) Full range XRD patterns of TiN and (Ti,Mg)N coatings, (b) Detailed view of XRD patterns around diffraction peak of (111) crystal plane.
Fig 1. Schematic of the coating system.
valuable insights on the effect of Mg ion concentration on HA formation and properties, high magnesium containing solutions fail to simulate the actual body environments so could not be used to comment on HA formation in vivo conditions. Relying on the positive role of Mg2+ ions on osteo-integration of implants and growth kinetics of HA, two approaches can be used for the increased availability of Mg2+ ions on implant surfaces: 1. Using Mg containing alloys as implant materials that have ability to release Mg in a controlled manner into the body fluid. However, pure magnesium and Mg based alloys are not suitable materials for implants since they have low corrosion resistance in most electrolyte mediums [18,20] limiting their usage in the human body environment. 2. Deposition of bio–compatible coatings containing a desired amount of magnesium on implant materials to give the implant surface a “magnesium releasing ability” in a controlled manner. The methodology and hypothesis of this study is based on the second approach; namely modifying the implant surfaces with magnesium containing coating that will contribute to HA formation on the implant surfaces in the body environment. For this purpose, a biocompatible coating that can be alloyed with magnesium is required. As the biocompatible coating, TiN, which is a widely used coating in biomedical applications, was selected. In order to dope TiN with magnesium, cathodic arc physical vapor deposition technique was utilized. In the literature there are few studies [55,56] on magnesium containing TiN coatings and these studies were mainly focused on optical [55] and high temperature oxidation behaviors [56]. These studies showed that magnesium atoms substituted the titanium atoms in
Table 1 Chemical Composition of 1X and 5X simulated body fluids. 1X SBF
b)
5X SBF
Order
Reagent
Amount (gr) in 40 ml solution
Reagent
Amount (gr) in 40 ml solution
#1 #2 #3 #4 #5 #6 #7 #8
NaCl NaHCO3 KCl K2HPO4 .3H2O MgCl2.6H2O HCl (1 M) CaCl2 Na2SO4
0.319 0.013 0.009 0.009 0.01 1.48 ml 0.01 0.003
CaCl2 MgCl2.6H2O NaHCO3 K2HPO4 .3H2O Na2SO4 KCl NaCl HCl (1 M)
0.056 0.06 0.07 0.045 0.015 0.045 1.6 Used for adjusting pH to 6
#9
(CH2OH)3CNH2
0.24
the cubic TiN structure forming (Ti1−x,Mgx)N phase even at high magnesium concentrations (x = 0.67). In the present study, the role of magnesium released from (Ti,Mg) N coatings on nucleation and growth of HA in SBF was investigated. In order to compare and understand the effect of magnesium on HA formation, magnesium free TiN coatings were also deposited and tested in two different SBF solutions with or without magnesium ions. The surface of the coatings and the HA films formed in SBF were characterized by SEM, XRD and FTIR techniques. 2. Material and methods 2.1. Deposition and characterization of (Ti,Mg)N and TiN coatings The coatings were deposited on grade 2 Ti plates (15x15x1 mm) by using cathodic arc physical vapor deposition technique. Our previous results [57] on the corrosion resistance of (Ti1–x,Mgx)N revealed that coatings with more than x = 0.12 Mg were not stable in chloride containing environments and better corrosion resistance has been achieved with x = 0.057. The parameters of the deposition process were, thus, selected to produce coatings containing x ≈ 0.06 Mg. Before the deposition, the substrates were polished using 1000, 2500 and 4000-grit SiC papers sequentially. Shortly after the polishing, the substrates were ultrasonically cleaned in acetone and ethanol baths for 10 minutes. After cleaning, the substrates were placed in the vacuum chamber then the chamber was evacuated down to 5 × 10−3 Pa. Since evaporation rate of Mg is higher than that of Ti, two Ti and one Mg cathodes were used to adjust Mg content of the coatings. The arc currents were 75 A for each Ti cathodes and 40 A for the Mg cathode. Before the coating process, the substrate surfaces were heated and etched by Ti ions produced as a result of cathodic arc evaporation. In the ion heating/etching process, a series of high bias voltages were applied on the substrates (–600, –800 and –1000 V in 30 sec for each step). The deposition process was conducted at 1 Pa N2 pressure and at –250 V bias voltages. In order to produce homogeneous coatings, the samples were rotated so that they could be exposed to each cathode for equal time intervals. Total coating time was 1 hour for each coating. The schematic of the coating system was given in Fig. 1. In order to understand the role of Mg on HA formation, a magnesium free TiN coating was also deposited. The magnesium content of the coatings was analyzed by using scanning electron microscope (SEM) with energy dispersive X-ray spectrometer (EDS) (JEOL JSM 7000F). The phase structure of the coatings and the HA deposits were analyzed by using glancing angle X-ray diffractometer with thin film attachment (Philips Model
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Fig. 3. SEM micrographs of HA deposits on TiN and (Ti,Mg)N after 4 weeks in 1X SBF.
PW3710). The Cu–Kα radiation and “θ scan method” with fixed incidence angle of 0.5° were used. After SBF experiments, the deposits formed on the coatings were investigated by using FTIR spectrometer and SEM. 2.2. Formation of hydroxyapatite in simulated body fluid Two different simulated body fluid concentrations (1X SBF and 5X SBF) were prepared as described in the references [58] and [59] (Table 1). The final pH values of 1X SBF and 5X SBF were adjusted to 7.4 and 6.5, respectively at 37 °C by using 1 M hydrochloric acid
or 1 M sodium hydroxide. The samples were incubated in 1X SBF for 1, 2, 3 and 4 weeks in order to monitor the HA formation. In experiments conducted in 5X SBF, HA film formation was accelerated so the samples were kept in the solution for only 3 days. TiN and (Ti,Mg)N coated samples were also tested for 3 days in magnesium free 5X SBF solution to investigate the role of Mg on the kinetics of HA formation. The HA deposits formed on TiN and (Ti,Mg)N coated samples in 1X, 5X and Mg free 5X SBF solutions were characterized by using FTIR spectroscopy and XRD techniques. 3. Results and discussions 3.1. Characterization of (Ti,Mg)N and TiN coatings
TiN peak (111)
Before SBF
TiN
(Ti,Mg)N Counts (a.u)
After SBF
The thickness of the coatings as determined from their fracture cross sections were 1.1 μm and 1.2 μm for TiN and (Ti,Mg)N, respectively. The EDS results showed that the (Ti1−x,Mgx)N coating contained Mg equivalent to x = 0.064. In the X-ray diffraction patterns of both coatings, characteristic TiN peaks at 36.37° (111), 42.43 (200), 61.63° (220) and 73.85° (311) were observed as shown in Fig. 2a. Furthermore, the peaks corresponding to magnesium or magnesium nitride phases were not present in the XRD pattern of the (Ti,Mg)N sample. Lattice parameters of the TiN and (Ti,Mg)N coatings were calculated from the XRD patterns by using “Philips X'PERT Plus Crystallography and Rietveld Analysis Software”. The calculated lattice parameters of the TiN and (Ti,Mg)N coatings were a = 4.248 Å and a = 4.253 Å, respectively. Magnesium addition shifted the TiN peaks to lower diffraction angles (Fig. 2b) expanding
1 week
2 weeks
3 weeks 4 weeks 35
36
37
38
2 Theta(deg) Fig. 4. XRD patterns of (Ti,Mg)N coatings after 1X SBF test.
Fig. 5. FTIR spectra of deposits formed on TiN and (Ti,Mg)N coatings in 1X SBF.
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Fig. 6. SEM micrographs of (a) (Ti,Mg)N and (b) TiN coatings after 5X SBF tests.
In 1X SBF, HA formation on the (Ti,Mg)N coating was initiated after the first week of immersion. HA deposits on the TiN coating, on the other hand, reached to an observable size only after 4th week. Higher amount of HA deposits on the (Ti,Mg)N could clearly been observed at the end of 4th week in the SEM micrograph (Fig. 3). The HA deposits, however, did not seem to homogeneously cover the whole surface after 4 weeks and the HA derived peaks were not observed in the XRD patterns of both coatings. Nevertheless, the strongest (111) TiN peak of the (Ti,Mg)N coating shifted to the diffraction angle of magnesium free TiN (111) peak after 1st week of immersion (Fig. 4). The shift of TiN peaks could be associated with the dissolution of Mg from the (Ti,Mg)N structure. After the first week, the FTIR spectrum (Fig. 5) of the (Ti,Mg)N sample exhibited PO−3 peaks around 553 and 4 1021 cm−1. Although PO−3 peaks in the spectrum of TiN started to 4 evolve after second week, their intensity were not comparable to those of the (Ti,Mg)N even after 4 weeks. Since the HA deposits formed in 1X SBF did not continuously cover the surfaces after 4 weeks, the HA formation experiments were
repeated in 5X SBF solution for a duration of 3 days. The SEM results showed that incubation of the samples in this solution led to a complete coverage of both TiN and (Ti,Mg)N surfaces with HA deposits after 3 days (Fig. 6). It could be clearly seen that the HA nodules on the (Ti,Mg)N surface developed to bigger sizes compared to those on the TiN surface (i.e. 2–3 times larger). The XRD patterns of the HA deposits on both coatings exhibited characteristic HA peaks (Fig. 7). FTIR spectra of the HA deposits on both coatings were given in Fig. 8. In the spectra, PO−3 group peaks at 475, 561, 600, 958, 1014 and 4 1115 cm−1, HPO−2 group peak at 867 cm−1, CO−2 peak at 1571 cm−1 4 3 −1 and H2O peak at 1634 cm are characteristic peaks of a hydroxyapatite structure. These results were in accordance with the literature; Suchanek et al. [21] and Hanifi et al. [60] reported that PO−3 peaks be4 came smaller with increasing amount of Mg in HA. Suchanek also reported that HPO−2 peak at 870 cm−1 evolved with higher magne4 sium contents in the HA [21]. In our study, PO−3 peak at 475 cm−1 4 −1 disappeared and CO−2 peaks at 1458 and 1571 cm became more 3 pronounced in the spectrum of the (Ti,Mg)N sample (Fig. 8a) when compared with the corresponding peaks of the TiN sample (Fig. 8b). Hanifi et al., and Suchanek et al. also reported that substitution of CO−2 3 for PO−3 was particularly observed in the Mg substituted HA structure. 4 −3 The similarities of CO−2 and HPO−2 peaks with the existing 3 , PO4 4 literature [21,60] is a strong indication of the Mg substituted HA formation on the (Ti,Mg)N coated sample. It could therefore be stated that the
Fig. 7. XRD analysis of TiN and (Ti,Mg)N coatings after 5X SBF tests.
Fig. 8. FTIR spectra of (a) (Ti,Mg)N and (b) TiN coatings after 5X SBF tests.
the TiN lattice by 0.17 %. Hodroj et al. [56] and Fenker et al. [55] reported similar results showing that magnesium substitution for titanium atom in TiN structure resulted in lattice expansion of TiN phase because of higher atomic radius of magnesium. 3.2. Hydroxyapatite formation in Simulated Body Fluids
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Fig. 9. SEM micrographs of (a) (Ti,Mg)N and (b) TiN coatings after Mg-free 5X SBF tests.
incorporation of Mg into the TiN structure promoted the Mg substituted hydroxyapatite formation on the coating. For revealing the role of Mg2+ ions on the growth of HA for both coatings, the 5X SBF tests were repeated in the SBF without Mg2+ ions. After 3 days of incubation, there was no HA formation on the TiN sample (Fig. 9b). Barrere et al. [54] showed the key role of Mg2+ ion on the formation of HA on Ti6Al4V substrates. They immersed the Ti6Al4V substrates in 5X SBF with and without Mg2+ ions. After 24 hours of immersion, they observed that HA deposits could only be formed in the solution containing Mg2+ ions. The results of our experiments further verified the importance of magnesium presence in solution during HA formation, this time in TiN coated samples. It should, however, be noted that HA formation was observed on the surface of (Ti,Mg)N coating in 5X SBF solution even in the absence of Mg2+ ions (SEM micrographs, Fig 9a). The HA derived peaks in the FTIR spectrum (Fig. 10) of the (Ti,Mg)N sample also showed the HA formation on the (Ti,Mg)N coating in this solution. These results clearly showed that the Mg in the (Ti,Mg)N coating promoted HA formation by releasing magnesium ions to the “SBF/ coating” interface.
cathodic arc PVD technique. The XRD patterns of the (Ti1–x,Mgx)N coatings indicated that addition of Mg equivalent to x = 0.064 did not form a separate Mg or magnesium nitride phases, instead, Mg incorporated in the TiN structure shifting the TiN diffraction peaks in the XRD pattern. In consequence of Mg dissolution into the simulated body fluid during the HA formation process, Mg-substituted HA formation was facilitated. In other words, the Mg doped TiN structure acted as a magnesium source promoting the Mg-substituted HA formation in simulated body fluids. This encouraging result offers the potential to produce hard tissue implants with the ability to induce on-site HA production on the surface, allowing improved tissue integration.
4. Conclusion
References
The effect of Mg on HA formation on TiN and (Ti,Mg)N coatings was studied. The coatings were deposited on Ti substrates by using
[1] B. Leukers, H. Gülkan, S.H. Irsen, S. Milz, C. Tille, M. Schieker, H. Seitz, Hydroxyapatite scaffolds for bone tissue engineering made by 3D printing, J. Mater. Sci. Mater. Med. 16 (2005) 1121–1124. [2] S. Teixeira, M.A. Rodriguez, P. Pena, A.H. De Aza, S. De Aza, M.P. Ferraz, F.J. Monteiro, Physical characterization of hydroxyapatite porous scaffolds for tissue engineering, Mater. Sci. Eng. C 29 (2009) 1510–1514. [3] P.K. Bajpai, C.M. Fuchs, Development of a hydroxyapatite bone grout, in: C.W. (Ed.), Proceedings of the First Annual Scientific Session of the Academy of Surgical Research, Pergamon Press, NY, 1985. [4] T.J. Webster, R.W. Siegel, R. Bizios, Bioceramics, World Scientific, New York, 1998. [5] T.J. Webster, R.W. Siegel, R. Bizios, Osteoblast adhesion on nanophase ceramics, Biomaterials 20 (1999) 1222–1227. [6] Y. Liu, T. Jiang, Y. Zhou, Z. Zhang, Z. Wang, H. Tong, X. Shen, Y. Wang, Evaluation of the attachment, proliferation, and differentiation of osteoblast on a calcium carbonate coating on titanium surface, Mater. Sci. Eng. C 31 (2011) 1055–1061. [7] W.K. Lee, S.M. Lee, H.M. Kim, Effect of surface morphology of calcium phosphate on osteoblast-like HOS cell responses, J. Ind. Eng. Chem. 15 (2009) 5. [8] J. Huang, S.M. Best, W. Bonfield, T. Buckland, Development and characterization of titanium- containing hydroxyapatite for medical applications, Acta Biomater. (2010) 241–249. [9] N. Sato, K. Kubo, M. Yamada, N. Hori, T. Suzuki, H. Maeda, T. Ogawa, Osteoblast mechanoresponses on Ti with different surface topographies, J. Dent. Res. 88 (2009) 812–816. [10] I. Wall, N. Donos, K. Carlqvist, F. Jones, P. Brett, Modified titanium surfaces promote accelerated osteogenic differentiation of mesenchymal stromal cells in vitro, Bone 45 (2009) 17–26. [11] K.Y. Cai, Y. Hu, K.D. Jandt, Y.L. Wang, Surface modification of titanium thin film with chitosan via electrostatic self-assembly technique and its influence on osteoblast growth behavior, J. Mater. Sci. Mater. Med. 19 (2008) 499–506. [12] M.D. Ball, S. Downes, C.A. Scotchford, E.N. Antonov, V.N. Bagratashvili, V.K. Popov, W.J. Lo, D.M. Grant, S.M. Howdle, Osteoblast growth on titanium foils coated with hydroxyapatite by pulsed laser ablation, Biomaterials 22 (2001) 337–347. [13] F. Ren, et al., Synthesis, characterization and ab initio simulation of magnesium substituted hydroxyapatite, Acta Biomater. 6 (2010) 2787–2796.
Fig. 10. FTIR spectra of (a) TiN and (b) (Ti,Mg)N coatings after Mg-free 5X SBF tests.
Acknowledgments This study was supported by The Scientific and Research Council of Turkey (TUBITAK –Project #112M339). The authors gratefully thank Mehmet Dokur for his valuable helps in FTIR analyses and ITU Graduate School of Science, Engineering and Technology for ITU-BAP Project #34525.
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[14] B. Bracci, et al., Effect Of Mg(2+), Sr(2+), And Mn(2+) on the chemico–physical and in vitro biological properties of calcium phosphate biomimetic coatings, J. Inorg. Biochem. 103 (2009) 1666–1674. [15] S.D. Garcia, D. Laurencin, J.M. Krafft, S. Casale, M.E. Smith, H.L. Pernot, G. Costentin, Influence of magnesium substitution on the basic properties of hydroxyapatites, J. Phys. Chem. C 115 (2011) 24317–24327. [16] A. Gozalian, A. Behnamghader, M. Daliri, A. Moshkforoush, Synthesis and thermal behavior of Mg-doped calcium phosphate nanopowders via the sol gel method, Sci. Iran 6 (2011) 1614–1622. [17] I. Cacciotti, et al., Mg-substituted hydroxyapatite nanopowders: synthesis, thermal stability and sintering behavior, J. Eur. Ceram. Soc. 14 (2009) 2969–2978. [18] L. Li, J. Gao, Y. Wang, Evaluation of cyto-toxicity and corrosion behavior of alkali heat-treated magnesium in simulated body fluid, Surf. Coat. Technol. 185 (2004) 92–98. [19] J. Takaya, H. Higashino, Y. Kobayashi, Intracellular magnesium and insulin resistance, Magnes. Res. 17 (2004) 126–136. [20] M.P. Staiger, A.M. Pietak, J. Huadmai, G. Dias, Magnesium and its alloys as orthopedic biomaterials: a review, Biomaterials 27 (2006) 1728–1734. [21] W.L. Suchanek, K. Byrappa, P. Shuk, R.E. Riman, V.F. Janas, K.S. TenHuisen, Preparation of magnesium-substituted hydroxyapatite powders by the mechanochemical– hydrothermal method, Biomaterials 25 (2004) 4647–4657. [22] M. Toriyama, S. Kawamura, Synthesis of b-tricalcium phosphate by use of wet milling, J. Ceram. Soc. Jpn. 94 (1986) 1004–1007. [23] M. Toriyama, S. Kawamura, Sinterable powder of mechanochemically synthetic b-tricalcium phosphate, J. Ceram. Soc. Jpn. 95 (1987) 741–745. [24] M. Toriyama, S. Kawamura, Y. Ito, H. Nagae, Thermal change of calcium deficient apatite obtained by mechanochemical treatment, J. Ceram. Soc. Jpn. 97 (1989) 554–558. [25] M. Otsuka, Y. Matsuda, J. Hsu, J.L. Fox, W.I. Higuchi, Mechanochemical synthesis of bioactive material: effect of environmental conditions on the phase transformation of calcium phosphates during grinding, Biomed. Mater. Eng. 4 (1994) 357–362. [26] M. Toriyama, A. Ravaglioli, A. Krajewski, C. Galassi, E. Roncari, A. Piancastelli, Slip casting of mechanochemically synthesized hydroxyapatite, J. Mater. Sci. 30 (1995) 3216–3221. [27] M. Toriyama, A. Ravaglioli, A. Krajewski, G. Celotti, A. Piancastelli, Synthesis of hydroxyapatite-based powders by mechanochemical method and their sintering, J. Eur. Ceram. Soc. 16 (1996) 429–436. [28] Y. Yokogawa, M. Toriyama, Y. Kawamoto, T. Suzuki, K. Nishizawa, F. Nagata, M.R. Mucalo, Preparation of calcium strontium apatite through mechanochemical method, Chem. Lett. 1 (1996) 91–92. [29] J. Liao, K. Ono, G. Kanayama, T. Isobe, M. Senna, Preparation of tricalcium phosphate by applying grinding technique and choosing hydroxy-containing initial materials, Chem. Sustainable Dev. 6 (1998) 233–236. [30] I.V. Fadeev, et al., Synthesis and structure of magnesium-substituted hydroxyapatite, Inorg. Mater. 39 (2003) 947–950. [31] Y. Zhao, C. Chen, D. Wang, Surf. Rev. Lett. 12 (2005) 401–408. [32] E. Van der Wal, A.M. Vredenberg, P.J. Ter Brugge, J.G.C. Wolke, J.A. Jansen, Biomaterials 27 (2006) 1333–1340. [33] M. Inagaki, Y. Yokogawa, T. Kameyama, Surf. Coat. Technol. 173 (2003) 1–8. [34] M. Roy, A. Bandyopadhyay, S. Bose, Induction plasma sprayed Sr and Mg doped nano hydroxyapatite coatings on Ti for bone implant, J. Biomed. Mater. Res. B 99B (2011) 2. [35] T. Huang, et al., Nanostructured Si, Mg, CO3 (2-) substituted hydroxyapatite coatings deposited by liquid precursor plasma spraying: synthesis and characterization, J. Therm. Spray Technol. 4 (2011) 829–836. [36] Y. Yunzhi, L.O. Kyo-Han Kim Joo, A review on calcium phosphate coatings produced using a sputtering process—an alternative to plasma spraying, Biomaterials 26 (2005) 327–337. [37] K.A. Gross, C.C. Berndt, Thermal processing of hydroxyapatite for coating production, J. Biomed. Mater. Res. 39 (1998) 580–587.
[38] K.A. Gross, C.C. Berndt, H. Herman, Amorphous phase formation in plasma-sprayed hydroxyapatite coatings, J. Biomed. Mater. Res. 39 (1998) 407–414. [39] H. Li, K.A. Khor, P. Cheang, Titanium dioxide reinforced hydroxyapatite coatings deposited by high velocity oxy-fuel (HVOF) spray, Biomaterials 23 (2002) 85–91. [40] Y. Yang, J.L. Ong, Bond strength, compositional and structural properties of HA coating on Ti substrate, ZrO2-coated Ti substrate, and TPS-coated Ti substrate, J. Biomed. Mater. Res. 64 (2003) 509–516. [41] S. Tao, H. Ji, C. Ding, Effect of vapor–flame treatment on plasma sprayed hydroxyapatite coatings, J. Biomed. Mater. Res. 52 (2000) 572–575. [42] X. Weichang, L. Xuanyong, Z. Xuebin, D. Chuanxian, Effect of hydroxyapatite coating crystallinity on dissolution and osseointegration in vivo, J. Biomed. Mater. Res. A 74A (2005) 553–561. [43] S. Ding, C. Ju, J.C. Lin, Characterization of hydroxyapatite and titanium coatings sputtered on Ti-6Al-4 V substrate, J. Biomed. Mater. Res. 44 (1999) 266–279. [44] C. Massaro, M.A. Baker, F. Cosentino, P.A. Ramires, S. Klose, E. Milella, Surface and biological evaluation of hydroxyapatite-based coatings on titanium deposited by different techniques, J. Biomed. Mater. Res. (Appl. Biomater.) 58 (2001) 651–657. [45] V. Nelea, C. Morosanu, M. Iliescu, I.N. Mihailescu, Hydroxyapatite thin films grown by pulsed laser deposition and radio-frequency magnetron sputtering: comparative study, Appl. Surf. Sci. 228 (2004) 346–356. [46] H. Zeng, W.R. Lacefield, The study of surface transformation of pulsed laser deposited hydroxyapatite coatings, J. Biomed. Mater. Res. 50 (2000) 239–247. [47] G. Jiang, D. Shi, Coating of hydroxyapatite on highly porous Al2O3 substrate for bone substitutes, J. Biomed. Mater. Res. (Appl. Biomater.) 43 (1998) 77–81. [48] Y. Han, K. Xu, J. Lu, Z. Wu, The structural characteristics and mechanical behaviors of nonstoichiometric apatite coatings sintered in air atmosphere, J. Biomed. Mater. Res. 45 (1999) 198–203. [49] Y. Han, T. Fu, J. Lu, K. Xu, Characterization and stability of hydroxyapatite coatings prepared by an electrodeposition and alkaline-treatment process, J. Biomed. Mater. Res. 54 (2001) 96–101. [50] L.A. De Sena, M.C. de Andrade, A.M. Rossi, G.D.A. Soares, Hydroxyapatite deposition by electrophoresis on titanium sheets with different surface finishing, J. Biomed. Mater. Res. 60 (2002) 1–7. [51] P. Habibovic, F. Barrere, C.A. van Bliterswijk, K. de Groot, P. Layrolle, Biomimetic hydroxyapatite coating on metal implants, J. Am. Ceram. Soc. 85 (2002) 517–522. [52] A.L. Oliveira, C. Elvira, R.L. Reis, B. Vazquez, J. San Roman, Surface modification tailors the characteristics of biomimetic nucleated on starch-based polymers, J. Mater. Sci. Mater. Med. 10 (1995) 827–835. [53] C. Ohtsuki, T. Kokubo, T. Yamamuro, Mechanism of HA formation of CaO-SiO2-P2O5 glasses in simulated body fluid, J. Non-Cryst. Solids 43 (1992) 84–92. [54] F. Barrere, et al., Nucleation of biomimetic Ca-P coatings on Ti6Al4V from a SBF x 5 solution: influence of magnesium, Biomaterials 23 (2002) 2211–2220. [55] M. Fenker, M. Balzer, H. Kappl, O. Banakh, Some properties of (Ti, Mg)N thin films deposited by reactive dc magnetron sputtering, Surf. Coat. Technol. 200 (2005) 227–231. [56] A. Hodroj, O. Chaix-Pluchery, P. Steyer, J.F. Pierson, Oxidation resistance of decorative (Ti,Mg)N coatings deposited by hybrid cathodic arc evaporation-magnetron sputtering process, Surf. Coat. Technol. (2011) 4547–4553. [57] B.S. Demirel, Iyon Demeti Destekli Manyetik Alanda Sıçratma Yöntemiyle (Ti,Mg) N İnce Filmlerin Üretilmesi, Yapısal ve Biyoaktivite Özelliklerinin Belirlenmesi, (MSc Thesis) Istanbul Technical University, Graduate School of Science, Engineering and Technology, Istanbul, 2003. [58] T. Kokubo, Apatite formation on surfaces o f ceramics, metals and polymers in body environment, Acta Biomater. 146 (1998) 2519–2527. [59] Y.F. Chou, et al., The effect of pH on the structural evolution of accelerated biomimetic apatite, Biomaterials 25 (2004) 5323–5331. [60] A. Hanifi, et al., Mg2 + substituted calcium phosphate nano particles synthesis for non viral gene delivery application, J. Mater. Sci. Mater. Med. 21 (2010) 2393–2401.