Biocompatibility and osteoconduction of active porous calcium–phosphate films on a novel Ti–3Zr–2Sn–3Mo–25Nb biomedical alloy

Colloids and Surfaces B: Biointerfaces 85 (2011) 103–115

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Review

Biocompatibility and osteoconduction of active porous calcium–phosphate films on a novel Ti–3Zr–2Sn–3Mo–25Nb biomedical alloy Sen Yu a,b,∗,1 , Zhentao Yu a,1 , Gui Wang b,1 , Jianye Han a,1 , Xiqun Ma a,1 , Matthew S. Dargusch b,1 a b

Northwest Institute for Nonferrous Metal Research, Xi’an, Shaanxi 710016, China CAST CRC, School of Mechanical and Mining Engineering, The University of Queensland, Brisbane, QLD 4072, Australia

a r t i c l e

i n f o

Article history: Received 6 September 2010 Received in revised form 17 January 2011 Accepted 16 February 2011 Available online 3 March 2011 Keywords: Porous calcium–phosphate films Osteoconduction Biocompatibility ˇ Titanium alloy Surface modification

a b s t r a c t The purpose of this study is to investigate the biocompatibility and osteoconduction of active porous calcium–phosphate films on the novel Ti–3Zr–2Sn–3Mo–25Nb biomedical alloy. The active porous calcium–phosphate films were prepared by the micro-arc oxidation method on the surface of a near ␤ biomedical Ti–3Zr–2Sn–3Mo–25Nb alloy, and then activated in a hydroxyl solution followed by an aminated solution. The phase composition, surface micro-topography and elemental characteristics of the active porous calcium–phosphate films were investigated with XRD, SEM, EDS and XPS. The biocompatibility was assessed using corrosion testing, the in vitro osteoblast cultivation test and implantation in soft tissue (subcutaneous and musculature). The osteoconduction was evaluated using the simulated body fluid test and by implantation in hard tissue. The results show that the active porous films are mainly composed of TiO2 anatase and rutile. The oxide layer is a kind of porous ceramic intermixture containing Ca and P. Immersion in simulated body fluid can induce apatite formation on the porous calcium–phosphate films resulting in excellent bioactivity. Cell cultures revealed that MC3T3-E1 cells grew on the surface exhibiting favorable morphologies. These results indicate that the Ti–3Zr–2Sn–3Mo–25Nb biomedical alloy coated with an active porous calcium–phosphate film has been shown to have excellent corrosion resistance, good biocompatibility and osteoconduction, which can promote cell proliferation and bone formation. Crown Copyright © 2011 Published by Elsevier B.V. All rights reserved.

Contents 1. 2.

3.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Materials and experimental methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1. Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2. Preparation and surface activation of the porous calcium–phosphate films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3. Characterization of the active porous films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4. Corrosion test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5. Cell culture test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.6. Fluorescence microscopy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.7. Evaluation of apatite-forming ability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.8. Implantation in soft tissue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.9. Implantation in hard tissue . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.10. Statistical analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Results and discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1. XRD and EDS analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2. Surface morphology of the micro-arc oxidation films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3. XPS analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

∗ Corresponding author at: Northwest Institute for Nonferrous Metal Research, Xi’an, Shaanxi 710016, China. Tel.: +86 29 86231084; fax: +86 29 86231084. E-mail address: [email protected] (S. Yu). 1 Present address: No. 96 Weiyang Road, Xi’an, Shaanxi 710016, China. 0927-7765/$ – see front matter. Crown Copyright © 2011 Published by Elsevier B.V. All rights reserved. doi:10.1016/j.colsurfb.2011.02.025

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3.4. Electrochemical tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5. Ca–P precipitation in SBF . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6. Osteoblast attachment and proliferation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7. Soft tissue reaction analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.8. Hard tissue reaction analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Discussion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction Titanium and titanium-based alloys are well established as suitable implant materials in the field of osteosynthesis, oral implantology, and in certain joint prosthetics applications. The attractiveness of titanium alloys for implantation is applications determined by a combination of favorable characteristics including corrosion resistance, biocompatibility, low elastic modulus, density and the capacity of titanium to bond with bone and other tissue [1–3]. Recent biomaterials research has focused on the development of ␤-titanium alloys because processing variables can be controlled to produce a set of properties to suit specific applications. The new alloys are being designed to have enhanced biocompatibility to promote acceptance of the implant by the body and reduced elastic modulus to promote load sharing between the implant and natural bone. New generation ␤ Ti-based alloy scaffolds are expected to combine high mechanical strength, low elastic modulus with diminishing risk of aseptic loosening due to stiffness mismatch (stress shielding). These alloys are designed without toxic alloying elements promoting good biocompatibility [3–9]. One recently developed promising biomedical alloy, Ti–3Zr–2Sn–3Mo–25Nb (TLM), shows significant improvement in these properties compared to previous generation titanium alloys such as CP titanium and Ti–6Al–4V. While the mechanical properties and corrosion resistance of Ti–6Al–4V are ideal for implant applications, studies have shown that both V and Al ions may cause long-term health problems. Moreover, the modulus of Ti–6Al–4V (110 GPa) is substantially higher than that of bone (10–40 GPa). Large modulus mismatches cause insufficient loading of bone adjacent to the implant (stress-shielding phenomena) and eventual failure of the implant. The newer TLM alloy, exhibit relatively low moduli while maintaining sufficient strength, and on the other hand, titanium (Ti), zirconium (Zr), stannum (Sn), molybdenum (Mo) and niobium (Nb) are believed to be non-toxic metals with good biocompatibility and they are widely used in different parts of the human body assistants in blood vessels, artificial valves in the heart, replacement implants in shoulders, knees, hips, elbows, ears and orodental structures [3,7–13]. However, being bioinert metallic materials, titanium and titanium-based alloys cannot bond to living bone directly after implantation into a host body. From a clinical point of view, the ideal biomaterial acting as a bone substitute should possess an osteoinductive ability, and biocompatibility [1,2]. Therefore, various surface modifications have been developed to improve the bioactive bone-bonding ability of titanium alloys. Of these, microarc oxidation (MAO) is one of the most extensively well-established methods, and the efficiency of this has been confirmed by many reports [1,2,14–18]. In this work, active porous Ca-, P-containing titanium oxide films were developed on the surface of the TLM alloy by the microarc oxidation method followed by activation treatment. The phase composition and surface characteristics of the films were investigated, and the biocompatibility and osteoconduction of the active porous films were tested using corrosion testing, the simulated

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body fluid test, the osteoblast cultivation test and in vivo experiments comparing the performance of the TLM alloy with the currently used Ti–6Al–4V(ELI) alloy. 2. Materials and experimental methods 2.1. Materials The compositions of the TLM and Ti–6Al–4V(ELI) alloys are shown in Table 1. Specimens for surface modification were prepared from plate with dimensions of Ø 10 mm × 4 mm. Samples for soft tissue implantations were Ø 1 mm × 10 mm, and those used for hard tissue implantation had the dimensions Ø 2 mm × 6 mm. The pre-treatment procedure included mechanical polishing with 2000 grit sandpaper, followed by etching in a mixture of hydrofluoric and nitric acids with volume fractions of 10% and 40%, respectively, for 5 s to remove the surface oxide, and then rinsing with ethanol followed by deionized water, and then dried. 2.2. Preparation and surface activation of the porous calcium–phosphate films The TLM alloy samples were used as anodes, and stainless steel plates were used as cathodes in an electrolytic bath. A fresh electrolyte containing 0.1 mol/L ␤-glycerophosphate disodium salt pentahydrate (C3 H7 Na2 O6 P 5H2 O, ␤-GP) and 0.9 mol/L calcium acetate monohydrate ((CH3 COO)2 Ca H2 O, CA) into deionized water was used for the MAO process. The MAO process was conducted at a fixed applied voltage in the range of 250–500 V using a direct current pulse power supply, and a pulse frequency, a duty circle, and a duration time set at 1000 Hz, 40% and 10 min, respectively. The system temperature was maintained below 40 ◦ C by a water bath during the anodizing process. After the above MAO treatment, the samples were washed with distilled water and dried in the drying cabinet. The samples obtained using the above MAO process were treated with an aqueous solution of potassium peroxydisulfate (100 g/L) for 8 h at 75 ◦ C dynamically and finally the samples were washed with hot water and dried in vacuum. Then the films were submitted to grafting reactions in an aqueous solution with the aid of the ceric ion technique using acrylamide (120 g/L) in nitric acid (0.04 mol/L) and ceric ammonium (0.04 mol/L) solution at 65 ◦ CC for 24 h under a stream of nitrogen. The samples were washed extensively with NaOH solution followed by hot water. 2.3. Characterization of the active porous films The phase and microstructure of the films on the samples were examined using X-ray diffractometry (CuK␣, scanning rate was 0.03◦ /s, scanning angle: 20–80◦ , the angle of the incident beam was fixed at 3◦ against the sample surfaces in order to detect the phase composition of the sample surfaces and the measurements were performed with a continuous scanning mode at a rate of 2◦ min−1 .). The microstructure of the films were observed using a scanning

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Table 1 The element composition of the Ti–3Zr–2Sn–3Mo–25Nb (TLM) and Ti–6Al–4V(ELI) alloys. Alloy Ti–3Zr–2Sn–3Mo–25Nb Ti–6Al–4V(ELI)

Nb

Mo

Sn

Zr

Al

V

C

N

O

25.1 –

2.90 –

2.02 –

3.08 –

– 5.8

– 4.05

0.01 0.012

0.03 0.031

0.14 0.09

electron microscope (JEOL, JSM-6460; JEOL Ltd., Tokyo, Japan) and an indication of the chemical composition of the films was obtained using energy dispersive X-ray spectroscopy (EDS) and energy dispersive X-ray analysis (AXIS ULTRA). 2.4. Corrosion test The electrochemical behavior of the TLM samples with and without surface modification was investigated using an IM6 electrochemical workstation (Zahner in Germany) in order to obtain the potentiodynamic polarization curves (E/I). The measurements were performed in a three-electrode flat cell with a Luggin capillary, a platinum slice as counter electrode, and a saturated calomel electrode was chosen as a reference electrode. The samples acted as the working electrode. Tyrode’s solution, which is commonly used for blood-contacting biomaterials, was used for the corrosion evaluation (deaerated with N2 at ambient temperature). The composition of Tyrode’s solution used was (g/L): NaCl(8.00), CaCl2 (0.20), KCl(0.20), NaHCO3 (1.00), MgCl2 (0.10), NaH2 PO4 (0.05). The polarization curves (E/I) were obtained by potentiodynamic scanning from 100 mV negative to the rest potential to 1.4 V at a rate of 0.1 mV/s, and the samples were first soaked in Tyrode’s solution for 30 min to measure the corrosion potential. 2.5. Cell culture test The biological properties of the samples were evaluated by preliminary in vitro cell tests. The MC3T3-E1 cells were used to characterize the proliferation and differentiation behavior of the cells. Osteoblasts are cultured upon the test substrates. The culture medium was 45% ␣-MEM (Gibco), 45% MEM (Gibco), 5% calf serum (Eurobio), 5% fetal bovine serum (Eurobio). The medium also contained gentamicin (50 ␮g/ml) and amphotericin-B (250 ␮g/ml). Cell proliferation is performed in 24-mutiwell plates, and 1 × 105 cm−2 growing cells contained in 1 ml culture medium are seeded in each well on the test samples and incubated at 37 ◦ C and 5% CO2 atmosphere with 95% humidity. 3 h enables complete adhesion of the cells on the test samples, the wells are carefully filled to 1 ml. The culture plate was transferred gently to a 37 ◦ C incubator. The incubation time durations were 2, 4 and 7 days. After each time point, the samples were taken out and rinsed with a phosphate-buffered saline solution (pH 7.2, PBS) twice to remove unattached cells. Cells grown on the test samples are fixed with sodium phosphate buffered with 2.5% glutaraldehyde. After two washes in the same buffer, the cells are postfixed with 1% OS O4 in saturated HgCl2 . After dehydration in graded ethanol, the cells were dried to the critical point, coated with gold and examined in a scanning electron microscope (JEOL JSM-6700F; JEOL Ltd., Tokyo, Japan). 2.6. Fluorescence microscopy After 2 days of culture, most of the samples were fixed in 2% paraformaldehyde/phosphate-buffered saline (PBS) at room temperature for 20 min, permeabilized in a PBS/Triton X-100 buffer (Sigma) (10 mM PBS, 0.2% Triton) for 15 min, and then blocked with 1% bovine serum albumin (BSA) (Sigma) in PBS (Sigma). Then, most of the samples were labeled with 1.2 mg ml−1 FITC-phalloidin (Sigma) for 60 min. After being washed in the buffer solution all the specimens were embedded in PBS/Glycerol-DABCO (1:1) (Sigma)

Ti Bal. Bal.

mounting medium and examined in a Nikon Eclipse-TE2000-U epifluorescence microscope (Nikon Corporation, Kanagawa, Japan). The others were fixed by paraformaldehyde for 20 min and rinsed with phosphate buffer solution three times. The MC3T3-E1 nuclei were then stained with Hoechst-33342 (1:1000, Sigma) according to the manufacturer’s instructions. 2.7. Evaluation of apatite-forming ability Simulated body fluid (SBF) was prepared by dissolving reagentgrade chemicals of NaCl, NaHCO3 , KCl, K2 HPO4 ·3H2 O, MgCl2 ·6H2 O, CaCl2 , and Na2 SO4 in distilled water and buffered at pH 7.4 with trishydroxymethyl-aminomethane ((CH2 OH)3 CNH2 ) and hydrochloric acid at 37 ◦ C, and the SBF with ion concentrations almost equal to those in human blood plasma. The sample was soaked in 50 ml of SBF in a plastic vial and stored in an oven at 37 ◦ C for 14 days, with the SBF being refreshed every day. After immersion for a pre-determined period of time, the samples were removed from the SBF, washed with distilled water and then dried. 2.8. Implantation in soft tissue For implantation in subcutaneous tissue, the implants were inserted into the dorsal bilateral subcutaneous tissue of 12 adult New Zealand rabbits (body weight > 3.0 kg), in a process approved by the local animal ethics committee (LAEC) at the Fourth Military Medical University, Xi’an. After the rabbits were anesthetized by intramuscularly injecting Ketamine (15 mg/kg) and diazepam (0.5 mg/kg), pentobarbital sodium (35 mg/kg; NEMBUTAL INJECTION, Dainabot, Osaka, Japan) was injected intravenously, and the implants were inserted in the dorsal subcutaneous tissue of the rabbits to observe the response of soft tissue. During the operation, the rabbits received an intravenous infusion of saline containing isepamicin sulphate for antibiotic. The operations were performed under the usual sterile conditions. For implantation in musculature, the implants were inserted into the dorsal muscles of 12 adult New Zealand rabbits (body weight > 3.0 kg) as above. Euthanasia was performed on the rabbits by using carbon dioxide after 2, 6, and 12 weeks of operation (n = 3 for each material and each time period), respectively. The soft tissue specimens were rinsed in sterile saline for 10 min and fixed in 10% neutral buffered formalin for 48 h at a volume more than 10 times that of the block section. After dehydration and embedded in paraffin. Paraffin embedded samples were sectioned at 30 ␮m thickness with a grinding system (EXAKT apparatus, Germany) and stained with Haematoxylin and Eosin (HE). These specimens were histopathologically observed with an optical microscope (VANOX-S, OLYMPUS, Tokyo, Japan). The number of inflammatory cells (neutrophil, Lymphocytes and plasmacell, macrophage and polykaryocyte) and fibroblast was assessed. The identification of inflammatory cells was based on their morphology. The thickness of fibrosa was also measured. 2.9. Implantation in hard tissue For hard tissue, the implants were inserted into the bilateral distal femora of 45 adult beagles (body weight > 3.5 kg). This process was also approved by the local animal ethics committee (LAEC) at the Fourth Military Medical University, Xi’an. Under anesthe-

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sia with Ketamine (15 mg/kg) and diazepam (0.5 mg/kg), the femur was exposed through a medical incision. Two holes, separated by about 10 mm apart, were carefully made in the lateral surface of the diaphysis of the femur using a low-speed dental round bur, with a physiological saline external coolant, and the implants were inserted press-fit into the metaphysic of the distal femora. The long axis of the cylinder was perpendicular to the long axis of the femur. Two TLM cylinders after MAO and activation treatment were implanted in one leg of the beagle and one bulk Ti–6Al–4V(ELI) and another TLM cylinder were implanted in the other leg as a paired control. The wound was then sutured. Before being killed, the beagles were given acheomycin (100 mg/kg/day) for 6 days. Thirty beagles were euthanized at 1, 2, 4, 8, and 12 weeks after operation (n = 6 for each material and each time period), respectively. Following euthanasia, bone specimens of 1, 2, 4, 8, and 12 weeks groups were separated into bone-implant blocks, and the tissue blocks were fixed in 10% neutral buffered formalin for 2 weeks at 4 ◦ C. After routine dehydration, embedding (Technovit 7200, Heraeus Kulzer, Germany) and slicing (EXAKT apparatus, Germany), the specimens were embedded in light-curing epoxy resin (Technovit 7200, Heraeus Kulzer, Germany) without decalcification. Embedded specimens were sawed perpendicular to the longitudinal axis of the cylindrical implants at a site 500 ␮m from the apical end of the implant. Specimens were ground to a thickness of 30 ␮m with a grinding system (EXAKT apparatus, Germany). Then the sections were stained with HE. In this way, histological sections of surrounding bone were obtained for tissue investigation with an optical microscope (VANOX-S, Olympus BX45, Tokyo, Japan). The percentage of bone in contact with the implants at the interface were quantitatively evaluated by an image analyzer (KS400, ZEISS, Germany) after tracing on a histological photograph. Thirty undecalcified specimens for each metal implant after1, 2, 4, 8, and 12 weeks were used for quantitative analysis (n = 6). The other fifteen beagles were euthanized at 1, 2, 4, 8, or 12 weeks after operation (n = 3 for each material and each time period), respectively. Following euthanasia, bone specimens of 1, 2, 4, 8, and 12 weeks groups were separated into bone-implant blocks, and the specimens were immediately immersed into pentadiol solution to fix the cell. The fixed specimens were split into two parts along the long axis of the implants. The samples were dehydrated, fixed with osmic acid again, and finally sputter coated with gold for examination under SEM (JEOL JSM-6700F; JEOL Ltd., Tokyo, Japan).

Fig. 1. XRD pattern of the TLM samples after surface modification treatment (a), and EDS spectra of the samples after treatment (b).

Fig. 1b shows the chemical compositions of the coating layers according to EDS. All the oxidized layers contained Ca and P as well as Ti and O irrespective of the presence of the Ca- and P-containing compounds. This implies that the elemental component in an electrolytic solution can be compounded into the films by using the MAO process [5,6,17,19,20], and it also implies that a small number of Ca- and P-containing phases are incorporated into the coating layer though less than the detection limit of XRD. 3.2. Surface morphology of the micro-arc oxidation films

3. Results and discussion

Fig. 2(a–d) shows SEM micrographs of the untreated and treated TLM alloy showing clearly the porous layer. The control surface had parallel grooves oriented along the polishing direction (Fig. 2a). The oxidized surface formed in the Ca- and P-containing electrolytic solution showed a porous microstructure where the micropores (∼2 ␮m in diameter) were distributed uniformly, as shown in Fig. 2b–d. These images show that the pores were relatively well separated and homogeneously distributed over the surface of the sample. During the early stage of the treatment, the formation of a barrier film was initiated by micro-arc oxidation in the electrolyte. By imposing a higher voltage, micro-arcing due to electrical breakdown of this dielectric layer became active and vigorous, resulting in disappearance of grooves and pore formation. A porous surface in implants is beneficial to bone tissue growth and enhances the anchorage of implants to the bone [1,6,21–23].

3.1. XRD and EDS analysis

3.3. XPS analysis

Fig. 1a shows XRD patterns of the crystalline phases of the coating layers formed by MAO and surface activation treatment. It can be seen that the oxidized layer was a mixture of TiO2 and a minor amount of amorphous phase with peaks of titanium. TiO2 was present in the form of anatase and rutile. No trace of Ca- and P-containing phases were detected by XRD.

Fig. 3 shows the XPS survey results of micro-arc oxidation films before and after surface modification. After surface activation of micro-arc oxidation films, elements of N appeared within the active functional group on the surface of the sample which had undergone surface activation. The analytical spectrum showed that the new nitrogen comes from an amino function group (399.8 eV). At the

2.10. Statistical analysis MinitabTM (Minitab Inc., USA) software was used for statistical analysis. The outliers were previously calculated using Grubb’s test and eliminated from the results. Normal distribution was determined using the Anderson–Darling test. The experimental data is expressed as the mean ± standard error deviation (SD). Statistical comparisons were carried out using analysis of variance (ANOVA, SAS Institute Inc., Cary, NC). A value of p < 0.05 was considered to be statistically significant.

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Fig. 3. XPS patterns of TLM alloy before and after surface modification treatment.

Fig. 4. Potentiodynamic polarization curves of samples with and without surface modification.

same time, no nitrogen was observed on the surface of the sample without surface activation. 3.4. Electrochemical tests Fig. 4 presents the polarization curves of the TLM alloy with and without surface modification treatment in Tyrode’s solution. Comparing the response to that of the unmodified TLM alloy, the anodic branch of the modified TLM alloy showed more passive behavior. Although the corrosion potential improved slightly after surface modification, the modified TLM alloy exhibited a larger passivation region (up to 600 mVSCE ) with relatively low current (about 20 times lower than the unmodified TLM alloy under the same potential). 3.5. Ca–P precipitation in SBF Fig. 2. SEM surface morphologies of the TLM titanium samples without treatment (a) and treated with MAO utilizing a Ca- and P-containing solution followed by activation (b): low magnification (×1000) (c) and high magnification (×5000) (d).

SEM images of the samples treated by MAO followed by bioactivation treatment by soaking in the SBF at 37 ◦ C for a period of time are shown in Fig. 5a–d. It is shown that a large amount of white scale-like particles are presented in some areas, such as pores and surface steps within 7 days in Fig. 5a. With an increase in soaking time up to 2 weeks, scale-like precipitates covered the whole porous surface of the samples, as shown in Fig. 5b. It is believed that the apatite nuclei were formed after approximately 14 days of

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Fig. 5. SEM images of the active MAO films soaked in SBF at 37 ◦ C for 7 days (a), 14 days (b), 28 days (c), 56 days (d) and the XRD pattern of the samples soaked in SBF at 37 ◦ C for 7 and 14 days (e).

immersion and continuing to grow, initially filling the pores, and then, spreading over the entire surface. By further increasing the SBF immersion time to 56 days, the entire surface of the active Ca–P porous coating was covered (Fig. 5c and d). The phase changes in the porous films during immersion in the SBF were investigated by XRD and are shown in Fig. 5e. Compared with Fig. 1a, distinct apatite peaks appeared after 14 days of immersion. All these results confirm that apatite was formed on the surface of the activated MAO films, and this indicates that the active Ca–P porous coating has good apatite-forming ability. 3.6. Osteoblast attachment and proliferation An in vitro test with MC3T3-E1 cells was used to examine the biocompatibility of the specimens produced using MAO followed by activation treatment. The morphologies of the MC3T3-E1 cells

grown on the porous specimens for 2, 4 and 7 days have been assessed by SEM and results are shown in Fig. 6. The osteoblastlike MC3T3-E1 cells grew well on the surfaces of the specimens after 2 days culture. Most of the cells were polygonally shaped, showing numerous, highly extended filipodia and very long pseudopods (Fig. 6a and b). In the latter stages of cell growth, the cells were in close contact with the specimen, with flatter surfaces, appeared to spread and there were a lot filopodia observed in the cell membrane (Fig. 6c and d). These filopodia act a storage area for extra cell membrane. Hence, as the cells attach to the surfaces, the excess membrane is consumed in the spreading process (Fig. 6d). In addition, lamellipodias and intercellular connections were clearly observed. At higher magnification (Fig. 6b), the pseudopods were fastened to the pores and steps on the surface, indicating an active cell migration and good livelihood. By the 2nd day, the cells already had begun to produce extra-cellular matrix on the surfaces (Fig. 6a)

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Fig. 6. SEM images of MC3T3-E1 cells cultured on the surface of active porous calcium–phosphate films for (a) 2 days, (b) 4 days, (c) and (d) 7 days.

Table 2 Soft tissue reaction to active porous calcium–phosphate implants at 2, 6 and 12 week time points (For implantation in musculature are shown in brackets). Time (weeks)

Thickness of fibrous membrane (␮m) (mean ± SD)

Inflammatory corpuscle Fibroblast

Neutrophil

Lymphocytes and plasmacell

Macrophage

Polykaryocyte

2

++ (++)

++ (+)

++ (++)

− (−)

+ (−)

13 ± 3.5* (12 ± 2.5)*

6

+++ (+++)

+ (+)

++ (+)

+ (+)

+ (−)

17 ± 3.0* (15 ± 2.5)*

12

+ (+)

− (−)

− (−)

− (−)

− (−)

12 ± 2.5* (10 ± 2.5)*

Gradings of cell density: (−) none; (+) a little; (++) moderate; (+++) a lot. Independent experiments: n = 3, statistical *p-value, *p < 0.05.

as well as forming a noticeable number of attachments between cells. This eventually led to complete confluence, where the surface was covered by a continuous sheet of cells (Fig. 6c). By the 7th day, all the surfaces showed evidence for the formation of an extra-cellular matrix (Fig. 6d) and confluence had been reached, except for small regions on the surfaces. This suggests that the prepared specimens provide a biocompatible environment which favours cell attachment. The attachment behavior of the cells was further characterized by immunofluorescence, as shown in Fig. 7. Cell attachment, proliferation and distribution is best viewed via fluorescence microscopy after staining with the nucleic acid dye, Hoechst 33342, as shown in Fig. 7a. The investigation showed that a number of cells were well attached to active porous surfaces and the cell density on the surfaces was significant within 2 days of culture, implying that the activated porous surfaces had favorable biological properties. The cytoskeletal organisation has been determined by actin labelling with FITC-phalloidin. The typical morphology of the MC3T3-E1 cells seen on the active porous calcium–phosphate films show that the cells are widespreaded over the surface, as illustrated in Fig. 7b and c. After 2 days of cell culture, cells appeared to be multipolar, the actin cytoskeleton was well organized for the cells on all the surfaces. Long bundles of assembled actin filaments, cross-

ing the whole cell were observed. The nucleus is well marked and the cytoplasm well granulated. The cells in contact with the other cells had an elongated spindle-like morphology. The cells appear to have generated offshoots which have contacted neighbouring cells. The staining demonstrates the presence of stress fibres in all cells grown on the sample. 3.7. Soft tissue reaction analysis Figs. 8 and 9 show the tissue response of the modified MAO TLM implants after placement in the subcutaneous tissue and musculature after 6 and 12 weeks, respectively. There was no significant injury to surrounding tissue (muscle, nerve, and blood vessels), and no appreciable inflammatory reaction to subcutaneous modified MAO implants in the adjacent soft tissue was observed (Fig. 8 and Fig. 9). For subcutaneous and intramuscular installation, a translucent extravascular fibrous membrane was found surround the inserted implants without conglutination at the early stage (Fig. 8a and Fig. 9a). At the 6-week time point, the inflammation consisted predominantly of neutrophils with some macrophages. A small quantity of lymphocytes was interspersed inside and adjacent to the membrane. Evidence of the proliferation of capillary vessels could be detected around the membrane, while no mate-

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Fig. 7. Adhesion of MC3T3-E1 cells on the active porous calcium–phosphate films after 2 days of cell culture: (a) fluorescence micrographs of pre-osteoblast nuclei stained with Hoechst 33342 illustrating cell attachment, (b) cell nuclei showed blue fluorescence after staining, (c) organization and assessment of actin cytoskeleton.

rial scrap and tissue degeneration were found (Fig. 8b and Fig. 9b). The thickness of the fibrous membrane in both tissues reduced, while the capillary vessel density increased for a culture time up to 12 weeks, but it thinner for musculature after 6 weeks (Table 2). In addition, the number of fibroblast, neutrophil, lymphocytes and plasmacell cells, and the macrophage and polykaryocyte density reduced with increasing culture time (Figs. 8c and 9c). These results

Fig. 8. The modified MAO sample inserted subcutaneously into a rabbit. Sections were stained with HE: (a) after 12 weeks, (b) after 6 weeks (original magnification ×200), (c) after 12 weeks (original magnification ×400).

suggest that the inflammation decreased substantially from the 6-week to the 12-week time point and was dominated by mononuclear cells such as lymphocytes and macrophages. The surrounding tissue showed evidence of healing, with the formation of granulation tissue and mild fibrosis over time.

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Fig. 9. The modified MAO sample inserted musculature into a rabbit. Sections were stained with HE: (a) after 12 weeks, (b) after 6 weeks (original magnification ×400), (c) after 12 weeks (original magnification ×200).

3.8. Hard tissue reaction analysis All dogs showed good tolerance to the surgical procedure. No animals exhibited infection at the surgical site, dislocation of the implants, or adverse reactions such as inflammation or foreignbody reactions on or around the implanted cylinders.

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For endosseous implants, histological characteristics were different for the modified TLM group, bare TLM group and bare Ti–6Al–4V(ELI) control group (Fig. 10). After week 2, bone tissue with an immature appearance had formed in an area that was distant from the implant surfaces in the Ti–6Al–4V(ELI), TLM and active porous TLM implants (Fig. 10b, g and i). But the bone around the untreated control appears to involve soft tissue that migrates in between the bone and implant surfaces, and a few lymphocytes were scattered within the soft tissue of the Ti–6Al–4V(ELI) group and TLM group (Fig. 10a and f). The active porous TLM implants are associated with vigorous bone formation that prevents soft tissue from intervening between the bone and implants, leading to direct bone deposition onto the implant surface (Fig. 10k and i). The lymphocytes were not observed at all times. Such differences in the implant morphogenesis of the interfacial bone are also clearly observed in the week 4, 8, 12 sections (Fig. 10c–e, h–j and m–o). Extensive spreading of the bone along the implant surface without soft tissue interposition is visible around the active porous TLM implants (Fig. 10m–o), whereas the bone around the untreated implants is largely separated from the implant surface by soft tissue (Fig. 10c–e, h–j). Another notable difference was the extent of intervention by soft tissue. The thickness of the soft tissue on the surface of the bulk Ti–6Al–4V(ELI) alloy is thicker than that observed on the TLM alloy for each time period. The surfaces of implants modified with the MAO followed by the activation treatment on the TLM alloy exhibit better biocompatibility and osteoconduction than the bulk Ti–6Al–4V(ELI) alloy and the untreated TLM alloy from the point of view of the quantity of inflammatory corpuscle and lymphocytes, and the amount of soft tissue around the implants for each time period. The performance of the TLM alloy without surface modification was better than that of the Ti–6Al–4V(ELI) alloy. Average histomorphometric values of bone-implant contact (Fig. 10) are shown (n = 6). Results are statistically significant when comparing the surfaces of the Ti–6Al–4V(ELI), and the TLM alloy with and without active porous surfaces. Fig. 11 shows the percentage of new bone in contact with the implant at the bone-metal interface after 1, 2, 4, 8 and 12 weeks. The percentage was around 6% after 1 week for the Ti–6Al–4V(ELI) alloy and the TLM implants while it was closer to 11% for the active porous TLM alloy. There were no significant differences between the Ti–6Al–4V(ELI) and TLM samples before the 2 week implantation. After 2 weeks, the percentage had markedly increased for each sample, and there were significant differences between the alloys (p < 0.05). Fig. 12 shows the surface morphologies of the implants after implantation for 1, 2, 4, 8, 12 weeks. In the first week, a little bulged cell, osteoblast, lamellar bone and collagen fibres can be observed on the bulk Ti–6Al–4V(ELI) and TLM implants (Fig. 12a). While in the active porous TLM group, a large number of osteoblast and Filipodia (woven bone) are present and a large number of collagen fibrous networks formed around the cells (Fig. 12k). In the following weeks (Fig. 12l–o), new born tissue increased on the active porous TLM plants, indicating that was their abundant growth of new bone on the surfaces of the active porous TLM alloy. After 12 weeks, almost all of the surface area on the activated porous TLM implants is directly covered with new bone, with no soft tissue found at the interfaces. The modified surface provide good biological fixation to the surrounding tissue through bone tissue ingrowth into the porous networks. There is no significant difference observed between the TLM and Ti–6Al–4V(ELI) alloys at the 1, 2 and 4 week time points. 4. Discussion The properties which are of interest for artificial bone substitution or bone-anchored reconstruction are corrosion resistance

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Fig. 10. New bone formation around Ti–6Al–4V(ELI), TLM and active porous TLM implants in femoral bone marrow of the femur. Sections were stained with HE: (a) Ti–6Al–4V(ELI), 1 w, (b) Ti–6Al–4V(ELI), 2 w, (c) Ti–6Al–4V(ELI), 4 w, (d) Ti–6Al–4V(ELI), 8 w, (e) Ti–6Al–4V(ELI), 12 w; and (f) TLM, 1 w, (g) TLM, 2 w, (h) TLM, 4 w, (i) TLM, 8 w, (j) TLM, 12 w; and (k) active porous TLM, 1 w, (l) active porous TLM, 2 w, (m) active porous TLM, 4 w, (n) active porous TLM, 8 w, (o) active porous TLM, 12 w. S: soft tissue, I: implant, B: bone.

in physiological environments, biocompatibility with living tissue after implantation into the bone, bioadhesion (bone ingrowth), modulus of elasticity, good processibility including joining and casting, and initial mechanical strength [1,3]. Materials with adequate pore structure have long been recognized as ideal bone substitutes because of good biocompatibility and osteoconductivity. However, the application of porous materials under load bearing conditions, such as in spinal interbody fusion and dental implants, have been restricted because of their poor mechanical properties [3]. Therefore, surface treatment to create porous surfaces on bulk Ti-based implants with appropriate mechanical properties is currently an important area in biomaterials research [1,3,6,20–25].

In the present study, the oxidized surface formed in the Ca- and P-containing electrolytic solution showed a porous microstructure with micron sized pores, and the porosity was interconnected in the films, as shown in Fig. 2b–d, which were relatively well separated and homogeneously distributed over the sample. The pore sizes in the MAO films can be expected to enhance cell proliferation (Fig. 6). The porous surface of the implants is beneficial to bone tissue growth and it also enhances the anchorage of the implants to the bone (Fig. 12), and a highly porous structure may be valuable as a depot for bioactive constituents such as growth factors or bone morphogenic protein. All metals and alloys are subjected to corrosion when in contact with body fluid as the environment within the body is very

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Fig. 11. The percentage of bone in contact with the implant after1, 2, 4, 8 and 12 weeks. Vertical lines show standard deviations. (*) Significant difference (n = 6, *p < 0.05).

aggressive as a result of the presence of water, dissolved oxygen, proteins, and various ions such as chloride and hydroxide. A variety of chemical reactions occur on the surface of a surgically implanted alloy. Implant corrosion caused by the reaction with body fluids and tissues seems to affect the fatigue life and ultimate strength of the material, leading to mechanical failure of the implants and releasing toxic metal ions to body. Further, the presence of particulate corrosion and wear products in the tissues surrounding the implant may result in a cascade of events leading to periprosthetic bone loss. Thus, corrosion resistance of a metallic implant material is consequently an important aspect of its biocompatibility. Although Ti based alloys exhibit reasonable to good corrosion resistance due to the formation of titania on their surfaces, the nature, composition and thickness of the protective oxide depends on the environmental conditions. Generally, degradation of the biomaterial surfaces occurs as a result of chemical reactions between titania and chloride ions in the media. It has been demonstrated that active porous calcium–phosphate films enhance the corrosion resistance of the substrate materials, as shown in Fig. 4. Potentiodynamic polarization curves show that the high corrosion resistance of the active porous calcium–phosphate films is mainly due to the barrier-type thickness (over ten microns) and the high degree of crystallinity of the film layer, which slows the mass transport process across the film to the substrate. EIS measurement is thought to be an effective way to analyze the corrosion behavior of coated metal substrates, so the active micro-arc oxidation films can enhance the thermodynamic stability and corrosion resistance of the TLM alloy. The film also has good anticorrosion properties that can reduce the corrosion rate of the matrix material. The corrosion resistance of the TLM alloy however is very good, which can be mainly attributed to the natural growth of the films on the surface of the TLM alloy, and the porous nature of the modified coating. The improvement in corrosion resistance of the coated TLM alloy is not very significant. During the micro arc discharge process under an applied voltage, intense plasma physical and chemical reactions occur in the discharge channel, accompanied by high pressure and high temperature conditions. Under these conditions, Ti near the surface of the titanium alloy substrate participates in the subsequent reactions and can be oxidized to form TiO2 . At the same time, Ca2+ , HPO4 2− or PO4 3− and OH− from the ionization of the CA, ␤-GP and H2 O move to the anode easily in the electrolyte under the applied electric field respectively and are incorporated into the coatings [1,5,6,20–25]. It has been demonstrated as shown in Fig. 1, that no trace of Ca- and P-containing phases were detected by XRD, which implies that a small number of Ca- and P-containing phases are incorporated into the coating layer though less than the detection

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limit of XRD, or the Ca- and P-containing phases were present as amorphous phases. It has been widely reported that the essential requirement for an artificial material to bond to living bone is the formation of a bone-like apatite layer on its surface in the body environment [22,24–26]. The current results reveal that the active porous calcium–phosphate coating has excellent induction capability for heterogeneous nucleation and growth of apatite in SBF. The in situ formation of apatite on the surface of the active porous TLM in SBF appears to be closely related to the Ca- and P-containing compounds. SBF is a metastable calcium phosphate solution supersaturated with respect to apatite. However, it has been reported that the barrier for the homogeneous nucleation of apatite is too high and a chemical stimulus is required in order to induce the heterogeneous nucleation of apatite from SBF [4,5,16,17,26]. The surface hydroxyl groups such as Ca(OH)2 , Ti(OH)4 , and COOH are known to be efficient inducers of apatite nucleation. In the porous films, CA was expected to undergo hydrolysis to form Ca2+ , OH− , and TiO(OH)2 in the SBF. Hydroxylated titanium oxide is considered to be insoluble and yields a TiOH surface, which might act as a nucleation site. The hydrolysis of the Ca- and P-containing phases provides Ca2+ , OH− , and HPO4 2− ionic species, which increase the local degree of supersaturation with respect to apatite near the surface. The provision of abundant OH− , NH2 − groups after activation treatment and the enrichment of calcium and phosphate trigger the nucleation of apatite on the oxidized TLM surface. When the apatite nuclei are formed, they spontaneously grow at the expense of calcium and phosphate ions from the metastable supersaturated SBF solution [5,6,15,17,25,26]. The well-established role that surface roughness and chemical composition play in determining the interaction of surgical implants with their surrounding cells and tissue and promoting clinical integration has led to the development of a diverse range of implants [21–23,27–34]. In the present study, the porous and rough surface produced by MAO increased cell attachment; this improvement was attributed to an increase in surface roughness [35]. The chemical composition of the surface layers also plays a crucial role in cell proliferation on the surface of the specimens. The incorporation of Ca and P ions into the oxidation process strongly encourages osteoblast differentiation. In this study, the incorporation of Ca and P-containing compounds into the oxide layer after MAO treatment supplied more Ca2+ and HPO4 2− or PO4 3− , and the OH− and NH2− active groups which have been introduced onto the surface of the films after the surface activation treatment. These ions bond to the negative ions of the protein resulting in a proportional increase in cell differentiation and propagation. In addition, the Ca2+ and PO4 3− and the OH− and NH2− active groups act as hydrophilic components in the surface oxides promoting surface hydration, which assists the proliferation and differentiation of the MC3T3-E1 cells [19,25,29–31]. Surface wettability may affect the attachment of cells either directly, since the attachment phase has emerged through an initial process involving physiochemical linkages between cells and surfaces including ionic forces or indirectly through alterations in the adsorption of conditioning molecules, e.g. proteins. Increased wettability enhances interaction between implant surfaces and the biological environment. Cell adhesion is generally better on hydrophilic surfaces [29–37]. The MAO treatment enables elements such as Ca and P to be incorporated into the surface oxide of the titanium alloys and the microtopography can be varied by regulating electrolyte and electrochemical conditions. The presence of Ca-ions has been reported to be advantageous for cell growth [5], and in vivo data shows that implant surfaces containing both Ca and P enhance bone apposition on the implant surface [25]. The present study confirms that the presence of Ca and P in the surface of titanium alloys plays an important role.

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Fig. 12. Scanning electron micrographs of the surface of implants at each stage after implantation, respectively. (a) Ti–6Al–4V(ELI), 1 w, (b) Ti–6Al–4V(ELI), 2 w, (c) Ti–6Al–4V(ELI), 4 w, (d) Ti–6Al–4V(ELI), 8 w, (e) Ti–6Al–4V(ELI), 12 w; and (f) TLM, 1 w, (g) TLM, 2 w, (h) TLM, 4 w, (i) TLM, 8 w, (j) TLM, 12 w; and (k) active porous TLM, 1 w, (l) active porous TLM, 2 w, (m) active porous TLM, 4 w, (n) active porous TLM, 8 w, (o) active porous TLM, 12 w.

5. Conclusion The biocompatibility and osteoconduction properties of the Ti–3Zr–2Sn–3Mo–25Nb alloy is better than that of the Ti–6Al–4V(ELI) alloy, the main reason is quite likely to be the effect of cytotoxic elements (Al and V). Micro-arc oxidation of the

TLM alloy in an electrolytic solution containing Ca- and P- subsequently produces activated porous titania on the surface. Apatite formation can be induced on these active porous titania layers by immersion in simulated body fluid. The active porous TLM surface offers an environment with a strong osteoblast-affinity, as demonstrated by the enhanced attachment, spreading, prolifera-

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tion and differentiation of the MC3T3-E1 cells. Subcutaneously and musculature test results revealed that there was no scrap material, tissue degeneration or appreciable inflammatory reaction to the subcutaneous modified MAO TLM implants in the adjacent soft tissue. Tests results from hard tissue implantation indicated that the TLM alloy produced with a porous and activated surface exhibits better biocompatibility and osteoconduction than unmodified Ti–6Al–4V(ELI) and untreated TLM. Micro-arc oxidation and surface activation pretreatment of TLM alloy surfaces markedly increased their osteoconductive capacity. New bone formation occurred extensively on the active porous TLM implants with virtually no intervention by soft tissue, maximizing bone-implant contact during healing up to nearly 100% at week 12. These results suggest that the TLM alloy with a surface modified to incorporate an active porous calcium–phosphate film has excellent corrosion resistance, good biocompatibility and osteoconduction, which can promote cell proliferation and bone formation. Acknowledgements The authors would like to acknowledge the financial support of Key Basic Research Program of China (2005CB623904), National Natural Science Foundation of China (30770586/C010515, 30870611/C100201), the Australia-China Special Fund, International Science Linkages Program co-supported by the Department of Innovation, Industry, Science and Research of Australia, and the Ministry of Science and Technology and National Science Foundation of China (31011120049), and the CAST CRC, which is established and supported by the Australian Government’s Cooperative Research Centers Program. The authors would also like to acknowledge Dr. Liying Zhang in the Fourth Military Medical University for donation of the MC3T3-E1 cell line. References [1] LiuF X., P.K. Chu, C. Ding, Surface modification of titanium, titanium alloys, and related materials for biomedical applications, Materials Science and Engineering R 47 (2004) 49–121. [2] D.F. Williams, On the mechanisms of biocompatibility, Biomaterials 29 (2008) 2941–2953. [3] M. Geetha, A.K. Singh, R. Asokamani, A.K. Gogia, Ti based biomaterials, the ultimate choice for orthopaedic implants—a review, Progress in Materials Science 54 (2009) 397–425. [4] D. Zaffe, Some considerations on biomaterials and bone, Micron 36 (2005) 583–592. [5] S.R. Paital, N.B. Dahotre, Calcium phosphate coatings for bio-implant applications: materials, performance factors, and methodologies, Materials Science and Engineering R 66 (2009) 1–70. [6] X. Li, C.A. van Blitterswijk, Q. Feng, F. Cui, Watari S F., The effect of calcium phosphate microstructure on bone-related cells in vitro, Biomaterials 29 (2008) 3306–3316. [7] Y. Okazaki, S. Rao, T. Tateishi, Y. Ito, Cytocompatibility of various metals and development of new titanium alloys for medical implants, Materials Science and Engineering A 243 (1998) 250–256. [8] M. Long, H.J. Rack, Titanium alloys in total joint replacement—a materials science perspective, Biomaterials 19 (1998) 1621–1639. [9] M. Niinomi, Recent research and development in titanium alloys for biomedical applications and healthcare goods, Science and Technology of Advanced Materials 4 (2003) 445–454. [10] D. Kent, G. Wang, Z. Yu, M.S. Dargusch, Pseudoelastic behaviour of a ␤ Ti–25Nb–3Zr–3Mo–2Sn alloy, Materials Science and Engineering A 527 (2009) 2246–2252. [11] Y. Zhentao, Z. Lian, Influence of martensitic transformation on mechanical compatibility of biomedical ␤ type titanium alloy TLM, Materials Science and Engineering A 438–440 (2006) 391–394.

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