Formation and growth of calcium phosphate on the surface of oxidized Ti-29Nb-13Ta-4.6Zr alloy

S. J. Li^a, R. Yang^{, , a}, M. Niinomi^b, Y. L. Hao^a and Y. Y. Cui<sup>a

a</sup> Institute of Metal Research, Chinese Academy of Sciences, 72 Wenhua Road, Shenyang 110016, PR China
^b Department of Production Systems Engineering, Toyohashi University of Technology, 1-1, Hibarigaoka, Tempaku-cho, Toyohashi 441-8580, Japan

Received 16 March 2003;  accepted 4 September 2003. ; Available online 18 November 2003.
Biomaterials
Volume 25, Issue 13 , June 2004, Pages 2525-2532

1. Abstract

The bioconductivity of a new biomedical titanium alloy Ti-29Nb-13Ta-4.6Zr achieved by a combination of surface oxidation and alkali treatment is reported in this paper. Oxidation treatment at 400°C for 24 h was found to result in the formation of a hard layer on the surface of the alloy. Immersion in a protein-free simulated body fluid and fast calcification solution led to the growth of calcium phosphate (Ca-P) phase on the oxidized and alkali-treated alloy, and the new bioconductive surface was still harder than the substrate. The surface processes during various treatment and immersion processes were investigated in detail, and the morphology of the calcium phosphate crystals was shown to be determined by the concentrations of Ca and P in the solution.

Author Keywords: Calcium phosphate coating; Crystal growth; Oxidation; Metal surface treatment; Titanium alloy

2. Article Outline

1. Introduction

2. Experimental

2.1. Materials and treatment

2.2. Immersion treatment

2.3. Surface characterization

3. Results

3.1. Characterization of oxidation treated TNTZ surface

3.2. Characterization of TNTZ surface after alkali treatment

3.3. Ca-P precipitation on the treated TNTZ

4. Discussion

5. Conclusion

Acknowledgements

References

3. 1. Introduction

Titanium alloys generally possess good mechanical properties and biochemical compatibility and are therefore preferred metallic materials for orthopaedic implant applications. As examples, Ti-6Al-4 V ELI, Ti-5Al-2.5Fe, and Ti-6Al-7Nb, have been developed for such purposes [1, 2 and 3]. However, none of these alloys can form a chemical bond with living bone. A common method to resolve this problem is physically forming a thin film of the highly biocompatible calcium phosphate (Ca-P) coating on the surface of the alloys. Processes for this purpose include dip coating [4], electron-beam deposition [5], hot isostatic pressing [6], pulsed laser deposition [7] and plasma spraying [8]. Of these methods, plasma spraying is most often used. Recently, new treatments that create a biocompatible layer mainly by chemical reactions after improvement of the bioconductivity of implant surfaces (without physically making a Ca-P thin film on the surface of the alloy) were developed [9, 10, 11, 12, 13 and 14]. Compared with previous physical methods, the chemical treatments are simpler, more economical and capable of producing uniform coating on complex shape implant.

Wear characteristics constitute another aspect of the performance of biomedical alloys. Failure generally occurs due to excessive wear of the components. The accumulation of wear debris will produce an adverse cellular response leading to inflammation, release of damaging enzymes, osteolysis, infection, implant loosening and pain [15 and 16]. The common approaches used recently to improve wear resistance of materials include surface modification, adjustment of alloy composition and heat treatment [17, 18, 19, 20 and 21].

A new alloy, Ti-29Nb-13Ta-4.6Zr, has recently been developed for biomedical application [22]. This alloy has low elastic modulus, high strength and toughness, and excellent corrosion resistance. Similar to other biomedical titanium alloys, the alloy also cannot form a chemical bond with bone directly. Our previous study shows that oxidation treatment of the alloy led to the formation of a hard layer on its surface which greatly improves its wear resistance [23]. If the surface after the oxidation treatment is bioconductive, a Ca-P layer will form therein after immersion in a biomimetic solution, and the wear resistance may be improved simultaneously. In this work, the surface characteristics after various treatments of preoxidized Ti-29Nb-13Ta-4.6Zr alloy, as well as the bioconductivity of the resultant surface, have been investigated.

4. 2. Experimental

2.1. Materials and treatment

A Ti-29Nb-13Ta-4.6Zr (mass%, the alloy is abbreviated as TNTZ below) ingot with a diameter of 60 mm was fabricated by induction skull melting using pure Ti, Nb, Ta and Zr as raw materials and then hot forged to rods of 20 mm in diameter. The composition of the experimental alloy obtained by wet chemical and gas analysis is 30.2 Nb, 12.4 Ta, 4.8 Zr, 0.16 O, 0.01 
(mass%) and balanced by Ti.

A section of the TNTZ ingot was heat treated at 790°C for 1 h under the protection of argon and then was divided into 10×10×2 mm³ plates by spark-erosion wire cutting. These plates were polished using SiC sand paper from 80 to 16 
m. The polished samples were ultrasonically cleaned in distilled water for 20 min, acetone for 20 min, 70% alcohol solution for 20 min, and then in distilled water again for 10 min. The cleaned samples were dried in air and used for the oxidation treatment and alkali treatment.

Table 1 shows that the surface of TNTZ alloy can be hardened significantly by oxidation treatment above 400°C for 24 h. The above surface oxidation treatment produces a simultaneous ageing effect on the bulk alloy, resulting in change of mechanical properties and Young's modulus. Because ageing at 400°C yields high strength and low modulus of the TNTZ alloy [24], the oxidation treatment in this study was carried out at 400°C in an electric furnace in favour of properties of the bulk alloy. Specifically, the cleaned samples were exposed in laboratory air at 400°C for 24 h (scheme TNTZ1 in Table 2). Some TNTZ1 samples were polished lightly to remove the oxide layer and to expose the solid-solution hardened metallic substrate, and these samples are referred to as TNTZ1^*.

Table 1. Vickers hardness of Ti-29Nb-13Ta-4.6Zr oxidation treated at different temperature for 24 h, tested under different load

Table 2. Schemes of heat treatment, alkali treatment, and soaking in biomimetic solutions

Note: (A) oxidized at 400°C for 24 h; (B) immersed in 10 
NaOH at 60°C for 24 h; (C) immersed in 10 
NaOH at 60°C for 24 h, and then heat treated in air at 600°C for 1 h; (D) immersed in SBF for 4 weeks; (E) immersed in FCS for 2 weeks. For each scheme, 3 samples were prepared and tested.

The oxidation treated samples were then alkali treated using a method developed by Kim et al. [9]. The oxidized samples were soaked in a 10
NaOH aqueous solution at 60°C for 24 h (scheme TNTZ2 in Table 2). For comparison, the TNTZ samples that were not oxidation treated were also alkali treated using this method (scheme TNTZ3 in Table 2). Then the alkali-treated substrates were gently washed with distilled water, and dried at 40°C for 24 h in an air atmosphere. Some TNTZ3 samples that were lightly polished to remove the new layer on their surfaces (for hardness test) were designated as TNTZ3^*.

2.2. Immersion treatment

A protein-free simulated body fluid (SBF) with ion concentrations nearly equal to those of human blood plasma [9] and a fast calcification solution (FCS) [10] were used for in vitro bioconductivity testing. The inorganic ion concentrations of the two solutions are listed in Table 3. The immersion experiments were carried out using the following procedure: each sample was immersed in 10 ml SBF (scheme TNTZ4 in Table 2) or FCS (scheme TNTZ5 in Table 2) in a polystyrene vial at 37°C for 2-4 weeks. The samples after immersion were thoroughly washed with distilled water and then dried in an oven at 40°C.

Table 3. Ion concentrations (m
) of blood plasma [9], SBF [9] and FCS [10]

2.3. Surface characterization

The surfaces of the samples after soaking in SBF or FCS were examined with a JSM-6301F scanning electron microscope (SEM), an EPM-810Q electron probe microanalyzer (EPMA), and a JAMP-7800 Auger electron spectrometer (AES). The samples were also subjected to X-ray photoelectron spectroscopy (XPS) analysis in a vacuum chamber at a base pressure below 3.98×10⁻⁵ Pa. In XPS analysis, an unavoidable 5 nm thick surface contamination layer was removed by sputtering with a cold cathode Ar+ ion gun (2.5 kV, 4 uA) at an angle of 50°. XPS spectra and the chemical shift of the photoelectron binding energy of each element detected were analysed by means of a 150 spherical sector analyser with a pass energy of 28 eV and a 240 W Al K
irradiation.

A Rigaku X-ray diffractometer adapted for thin-film (TF) analysis, with a RINT-2500 measuring system and Cu K
irradiation (18 kW HV generator at 40 kV per 250 mA), was used to examine phase constitutions of the surface layers of TNTZ after various treatments. Vickers hardness was measured using a hardness tester with a load of 10, 25 and 50 g held for 15 s.

5. 3. Results

3.1. Characterization of oxidation treated TNTZ surface

Fig. 1 shows an AES analysis of the depth profile of O, Nb, Ti, Ta and Zr in a TNTZ1 sample. It can be seen from this figure that the top layer (about 0.5 
m thick) of the sample contains a relatively high but constant concentration of oxygen, suggesting the existence of a distinct oxide layer. The concentration of oxygen decreases to the normal level of the bulk below the oxide layer. The morphology of the oxidized surface (Fig. 2b), however, was not obviously changed from the original surface (Fig. 2a). X-ray diffraction pattern presented in Fig. 3a shows that the main constituent of the surface film is rutile TiO₂ and no peaks due to Magnéli phases Ti_nO_2n−1 were observed. XPS results in Table 4 confirm that Ti is in its four-valent state, suggesting that the oxide is stoichiometric TiO₂ [25]. A little Nb₂O₅ and ZrO₂ were also detected by XPS (Table 4). The lattice spacing of the
-Ti phase of the oxidized layer was calculated from Fig. 3a to be greater than that of the normal
lattice of unoxidized samples. This lattice expansion is attributed to the presence of interstitial oxygen in the
solid solution.

(5K)

Fig. 1. AES analysis showing the depth profiles of Ti, O, Nb, Ta and Zr of TNTZ1.

(59K)

Fig. 2. SEM micrographs showing the surface morphology of (a) untreated TNTZ, (b) TNTZ1, (c) TNTZ3, and (d) TNTZ2 sample.

(8K)

Fig. 3. TF-XRD patterns of (a) TNTZ1, (b) TNTZ3, (c) TNTZ2, (d) TNTZ6, (e) TNTZ4, and (f) TNTZ5 sample.

Table 4. Binding energies of the main XPS peaks

Table 5 lists Vickers hardness values of TNTZ after oxidation treatment, nitrogen ion-implanted Ti-6Al-4 V and Ti-6Al-4 V oxidized at 400°C for 24 h, all tested under 10, 25, and 50 g load. These data reveal a significant increase in hardness caused by diffusion of oxygen into TNTZ. This effect is observed at load as high as 50 g which indicates an excellent penetration depth. For instance, the Vickers hardness of TNTZ1 under 25 g load was 607 MPa, significantly greater than that of nitrogen ion implanted or oxidized Ti-6Al-4 V. The hardened surface will be of great advantages to the improvement of wear resistance of TNTZ [23].

Table 5. Vickers hardness of TNTZ and Ti-6Al-4 V after various treatments, tested under different load

3.2. Characterization of TNTZ surface after alkali treatment

After alkali treatment, the oxidized surface became more porous and many cracks were found in the oxidation layer (Fig. 2d). TF-XRD pattern ( Fig. 3c) shows that an amorphous layer characterized by a halo at 2
of 20-35° was formed during the alkali treatment. Compared with the pattern of TNTZ3 (Fig. 3b), the intensity of the halo of TNTZ2 is much higher. XPS results given in Fig. 4 show that the main components of the TNTZ2 surface are Ti, O, and Na, with a little amount of Nb. Na was found to be in its ionic form Na⁺ and Ti is in its four-valent state, as confirmed by the Na 1 s signal at 1071.3 eV and Ti_2p^3/2 signal at 457.8 eV (Fig. 4 and Table 4). A typical O 1 s spectrum of the TNTZ2 surface layer (Fig. 4a) contains three oxygen species, O²⁻, OH⁻ and absorbed H₂O evidenced by the corresponding O 1 s peaks at 529.7, 531.6 and 533.0 eV, respectively (Table 4). These results suggest that oxygen in the surface layer exists in the form of oxide and hydroxide. The state of Na and Ti on the surface of TNTZ3 is identical with that on TNTZ2, but the amounts of Na⁺ are lower than that of TNTZ2. The TF-XRD and XPS results suggest that amorphous sodium titanate formed on the surface of TNTZ2 and TNTZ3. In order to confirm the composition of the sodium titanate, the sample was heat treated at 600°C for 1 h and furnace cooled to room temperature after oxidation and alkali treatments (scheme TNTZ6 in Table 2). The amorphous phase disappeared and large amounts of crystalline sodium titanate and TiO₂ precipitated (Fig. 3d). This experiment demonstrated that the amorphous phase was sodium titanate that can be crystallized by heat treatment [9].

(16K)

Fig. 4. XPS spectrum of (a) O 1 s, (b) Na 1 s, (c) Ti 2p, and (d) Nb 3d of TNTZ2 sample.

Some TNTZ2 samples were slightly polished and then tested for hardness. The results show that the surface hardness is still higher than that of the substrate, although it decreased to some degree.

3.3. Ca-P precipitation on the treated TNTZ

After immersion in SBF and FCS for 4 and 2 weeks, respectively, no precipitation was detected on oxidized TNTZ surface; however, some spherical and platy precipitations formed on oxidized and alkali-treated TNTZ surface, respectively. Immersion in SBF for 4 weeks yielded a uniform layer of Ca-P on the surface of TNTZ4 (Fig. 5). These precipitates are spherical and seem to be poorly crystallized as indicated by the high-magnification SEM micrograph of Fig. 5b. The EDX profiles of the cross-sections of TNTZ4 show that the concentrations of Ca and P gradually increase from the substrate toward the surface ( Fig. 5c), which indicates that the amorphous Ca-P layer formed during the present treatments was strongly bonded to substrate without forming a distinct interface. By contrast, after immersion in FCS for 2 weeks, well crystallized, platy Ca-P layers were produced on the surface of TNTZ5 ( Figs. 6a and b). These large, thin crystal plates had directly grown on the TNTZ surface (Fig. 6c). The main elements on the oxidized and alkali-treated TNTZ surface after immersion in SBF and FCS are Ca, P and O. This Ca-P layer is tentatively identified as hydroxyapatite (HA) because the binding energies of Ca at Ca_2p^3/2 and P at P_2p^3/2 are consistent with the standard data for HA (Table 4) [25]. The TF-XRD results shown in Figs. 3e and f confirmed the above deduction from XPS analysis: two typical peaks at 2
of 26° and 32° of apatite can be seen in the TF-XRD patterns of the precipitates deposited on TNTZ2 surfaces from SBF and FCS. The different morphology of the Ca-P layer on SBF-soaked from FCS-soaked TNTZ is most likely caused by the compositional difference between the two solutions.

(38K)

Fig. 5. Surface morphology of TNTZ4 sample showing (a) spherical deposition, (b) enlargement of (a), and (c) EDX profiles of a cross section.

(46K)

Fig. 6. Surface morphology of TNTZ5 sample showing (a) plate-like deposition, (b) enlargement of (a), and (c) plates directly grown from the substrate surface (edge-on view of the side of a specimen).

The growth of Ca-P coating on the unoxidized surface of titanium alloys was recently investigated [26]. In that study, pure titanium, Ti-6Al-4 V and Ti-29Nb-13Ta-4.6Zr were immersed in NaOH (10 
) for 24 h and heat treated at 600°C for 1 h. Then, these alkali-treated specimens were immersed in SBF to examine the Ca-P growth on the surface of the three alloys. The results indicated that Ca-P layer could easily form on the surface of pure titanium and Ti-6Al-4 V, whereas it was difficult to form on Ti-29Nb-13Ta-4.6Zr. The improvement of surface bioconductivity of Ti-29Nb-13Ta-4.6Zr by combined oxidation and alkali treatments may be due to the increased porosity and high Na concentration on the modified surface (see Section 3.2).

6. 4. Discussion

After oxidation treatment at 400°C for 24 h, an oxide layer containing primarily TiO₂ and small amounts of Nb₂O₅ and ZrO₂ (both can be used to improve the wear resistance of materials [27, 28, 29, 30 and 31]) formed on the surface of TNTZ. Beneath the oxide layer, there exists a diffusion-hardened layer which optimized the wear resistance of TNTZ. Compared with other chemical- or plasma-deposition methods aimed at improving the wear resistance of alloys, such as the formation of TiN, ZrO₂, SiC, and amorphous diamond-like carbon coating [32 and 33], the present method possesses great advantages because it provides a hard, non-porous and uniform surface with sufficient adhesion and cohesion strengths.

After immersion in SBF and FCS for 4 and 2 weeks, respectively, no Ca-P layer formed on oxidized TNTZ surface, indicating that the oxidized surface containing TiO₂ is not bioconductive enough to induce Ca-P precipitation from SBF and FCS. Additional treatments are needed to improve the bioconductivity of the oxidized surface. Many methods of improving the bioconductivity of implants have been developed [9, 10, 11, 12, 13 and 14]. Most of these methods attempt to produce TiO₂ in special forms, such as the porous structure, that can be negatively charged by OH⁻ groups in an aqueous environment. In this work, the alkali treatment developed by Kim et al. [9] was adopted. The alkali treatment results in the formation of an alkali titanate hydrogel layer on preoxidized TNTZ surface. When exposed to SBF or FCS, the alkali titanate layer is hydrated to transform into TiO₂ hydrogel via release of alkali ions from the alkali titanate layer into SBF or FCS, leading to the deposition of the Ca-P phase from the two solutions. It follows that by combining alkali treatment and oxidation treatment the TNTZ surface is made both harder and bioconductive, and a layer of Ca-P can be induced on this harder surface. The two-step surface modification method developed in this work is very promising for the application of TNTZ because its bioactivity and wear resistance have been improved simultaneously.

It should be noted that the morphology and deposition rate of Ca-P from SBF and FCS are different. The formation of HA can be viewed as a process of crystal growth from aqueous solution, so the above difference can be explained on the basis of second phase formation from a supersaturated solution. In this study, the concentrations of Ca are 2.5 and 3.1 m
in SBF and FCS, respectively, which is higher than the equilibrium saturation concentration of hydroxyapatite in aqueous solution under physiological condition [34]. Thus both solutions used in the present study can be regarded as supersaturated. According to the theory of solidification, the nucleation rate in aqueous solution, I, may be expressed as a function of the supersaturation,
, by a relationship of the form [35]

I=A exp[−K(log S)^−2],

where c is the concentration of the solution, c^* is the equilibrium saturation concentration at the same temperature, A and K are two constants depending on the physical properties and hydrodynamics of the system.

It can be deduced from the above equation that the nucleation rate I is determined by the supersaturation if other parameters are assumed unchanged. The supersaturation in FCS is higher than that in SBF, so the nucleation rate of HA in FCS is accordingly higher than that in SBF, which will result in faster formation of HA nuclei in FCS. The different morphology of crystals can also be explained by the difference in concentration between the two solutions. It is known from solidification theory that, in the course of solidification, crystal morphology and growth rate can be directly related to supercooling which may be produced by changes in temperature and/or composition. The higher the supercooling is, the faster the crystal grows. For the crystal growth in aqueous solution, the relationship between the supersaturation, S, and the supercooling, Δ
, can be characterized by the equation [35]

where c^* is the equilibrium saturation concentration and
the temperature.

The above equation shows that there exists a linear relationship between the supersaturation and supercooling. The concentrations of Ca and P in FCS are much higher than those in SBF (Table 3). Thus, the supercooling formed in front of growing HA in FCS is higher than that in SBF. Once the supercooled region forms, any projection that forms by chance on the liquid/solid interface will grow into the supercooled region rapidly, forming a dendritic or needle-like crystal. But for HA growth in SBF, because of the lower supercooling in front of the growing crystals, the solid/liquid interface will be stable and the crystal will grow slowly in the form of spherical structure to minimize surface energy. In a word, it is the difference in solute concentrations between the two solutions that results in the difference of morphology, size and growth rate of Ca-P crystals from SBF and FCS. Thus, it is possible to adjust the Ca-P microstructure of the coating on TNTZ surface to meet specific clinical requirement by selection of a suitable immersion solution. The results of hardness test show that, after alkali treatment, the hardness of the oxidized layer decreased to a certain degree. This means that the oxidized surface was partly deteriorated by treatment in the highly concentrated alkali solution. In order to maintain the excellent characteristics of the oxidized surface, methods that have much less adverse effect on substrate, such as that proposed by Wen et al. [10], may be used in the future.

7. 5. Conclusion

(1) Oxidation treatment at 400°C for 24 h led to the formation of a hard layer on Ti-29Nb-13Ta-4.6Zr surface. The oxides are mainly composed of TiO₂, with small amounts of Nb₂O₅ and ZrO₂; an oxygen diffusion layer exists beneath the surface oxide layer.

(2) A titanate layer forms on the preoxidized surface after alkali treatment, and growth of a layer of Ca-P has been successfully induced on the titanate layer by immersing in SBF or FCS solutions. The results show that a hard and bioconductive surface can be achieved on Ti-29Nb-13Ta-4.6Zr by combining the oxidation treatment and alkali treatment, leading to simultaneous improvement of its wear resistance and bioactivity.

(3) The difference in morphology of the Ca-P crystals grown from SBF and FCS solutions was interpreted in terms of phase transformation theory, and was attributed to the difference of the supersaturated concentrations of Ca and P in the two solutions.

8. Acknowledgements

The work was supported by the NSFC (grants 59925103 and 50828101), the Chinese MoST (grant TG2000067105), and NEDO and JSPS in Japan. SJL is grateful to the Ministry of Education, Science and Culture of Japan for the award of a visiting studentship.

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