In Vitro and in Vivo Degradation Behavior and Biocompatibility

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Bio-interactions and Biocompatibility

In vitro and in vivo Degradation Behavior and Biocompatibility evaluation of MAO-FHA coated Mg-Zn-Zr-Sr Alloy for Bone Application Yansong Wang, Minfang Chen, Xiao Li, Yun Zhao, Chen You, Yankun Li, and Guorui Chen ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.9b00564 • Publication Date (Web): 17 May 2019 Downloaded from http://pubs.acs.org on May 22, 2019

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ACS Biomaterials Science & Engineering

In vitro and in vivo Degradation Behavior and Biocompatibility evaluation of MAO-FHA coated Mg-Zn-Zr-Sr Alloy for Bone Application Yansong Wang†, Xiao Li†, Minfang Chen†, ‡,§*, Yun Zhao†, ‡,§, Chen You†,§, Yankun Li†, Guorui Chen† †School of Materials Science and Engineering, Tianjin University of Technology, Tianjin 300384, China ‡Tianjin Key Lab for photoelectric Materials & Devices, Tianjin 300384, China §Key Laboratory of Display Materials and Photoelectric Device (Ministry of Education), Tianjin 300384, China * Correspond Author email: [email protected]

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Abstract

Magnesium and its alloys are biodegradable materials with great potential for biomedical development, however their high rate of degradation in biological environments limit the widespread application of these materials. In order to improve the corrosion resistance of magnesium alloy, a functional calcium phosphate coating was prepared on Mg-3Zn-0.5Zr-0.5Sr alloy by micro-arc oxidation (MAO) combined with chemical deposition of fluoridated hydroxyapatite (FHA). A dense calcium-phosphorus coating 6 μm thick composed of needleshaped fluoridated hydroxyapatite formed on the surface of the MAO layer. The MAO-FHA coating exhibited good mineralization ability to induce hydroxyapatite deposition on its surface during degradation testing in simulated body fluid. Furthermore, the results of cytotoxicity tests and in vivo experiments revealed the MAO-FHA coating had good cell compatibility and promoted osteogenesis while decreasing inflammatory response, demonstrating the MAO-FHA coating was effective at enhancing the corrosion resistance and biocompatibility of magnesium alloy. Key word: magnesium; micro-arc oxidation; FHA; biocompatibility; cytotoxicity; in vivo Introduction Biodegradable materials are greatly favorable for patients with bone injuries because secondary surgeries following orthopedic implantation are avoided1-3. This reduces economic burden, physical distress and the possibility of various complications for patient4-7. Magnesium and its alloys are considered to be biodegradable materials with great development potential 8-11 especially for bone repair12. These biocompatible13-15 materials have densities similar to natural bone16-17 and mechanical properties showing no stress shielding effect18. However the 2 ACS Paragon Plus Environment

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degradation rates of magnesium and its alloys in the human body are fast19,20, which has limited the application of these materials21,22. Surface treatment methods used to improve the corrosion resistance of magnesium alloys include electrochemical deposition23, chemical conversion coating24,25, biodegradable polymer coating26-28, and micro-arc oxidation (MAO)29. The preparation of a bioactive functional coating on the surface of magnesium alloy, a major focus of magnesium alloy surface treatment research, has been shown to improve the corrosion resistance of magnesium alloy and its biological activity. Of the biomedical materials studied, calcium-phosphate ceramic coatings have been shown to be desirable for magnesium alloy surface treatments30. Hydroxyapatite (HA, Ca10(PO4) 6(OH)2) is a major component in natural bones and has a role in promoting bone growth. HA coating deposited on the surface of Mg alloy has been shown to be unstable and extremely prone to decomposing in bodily fluids31-35,39. HA doped with cations and anions such as Na+, Sr+, Mg2+, CO32–, and F– 36 has demonstrated improved stability. Among these, fluoridated hydroxyapatite (FHA, Ca10(PO4)6(OH)2-xFx, where 0 < x < 2)37 has been shown to degrade at a slower rate than HA coating38 and has good bioactivity31,41. The FHA coating also provides better apatite-like layer deposition, better cell adhesion, better protein adsorption, and improved alkaline phosphatase activity in cell culture32, F- ion of FHA can prevent the decrease of bone density caused by osteoporosis45,62, and also has an inhibitory effect on bacteria48,63. Therefore, FHA is a potential material for the functional coating of biodegradable magnesium alloy. In general, electro-deposition is used to prepare FHA coating on the surface of magnesium alloy. Bakhsheshi-Rad et al.32 electro-deposited FHA coating on the surface of Mg-Ca alloy and produced a dense, uniform FHA coating with a slow degradation rate and good apatite-inducing 3 ACS Paragon Plus Environment


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ability. Compared with Mg-Ca alloy, the rate of hydrogen evolution of the FHA coated alloy decreased from 4.98 ml/cm2/day to 0.92 ml/cm2/day. Additionally, the FHA coating was shown to significantly inhibit Mg-Ca alloy degradation. Li et al.40 also prepared FHA coating on the surface of Mg-Zn alloy using electrochemical deposition. In this study, human bone marrow stromal cells (hBMSCs) were utilized to investigate the cellular biocompatibility of Mg-6 wt.% Zn alloy after surface modification. The results demonstrated the bioactive FHA coating had more effective stimulating effects on hBMSCs proliferation and differentiation. Subsequently, the authors implanted a FHA-coated Mg-Zn alloy screw into an adult New Zealand rabbit femoral condyle41. The one month in vivo animal test showed enhanced interface bioactivity and more direct contacts around the FHA coated implant. Razavi et al.42 prepared MAO/FHA coatings on the surfaces of magnesium alloys by MAO and electrophoretic deposition techniques. Their results revealed a significant enhancement in the biocompatibility of MAO/FHA coated implants compared to uncoated implants. Implants with MAO/FHA coating also showed a decrease in plasma weight loss and released magnesium ion. According to histological results, new bone formation increased and inflammation decreased around the implant showing the use of FHA coating improved the biocompatibility of AZ91 magnesium alloy implants. Most studies have used electrochemical deposition42-44 to prepare FHA coatings on MAO coated Mg alloy. However, electrochemical deposition also has some inherent defects such as cracks and poor adhesion34,61. Compared with previous studies and methods, hydrothermal synthesis adopted in this study has the advantage of low cost45 and low energy consumption, simplicity, high purity, good dispersibility, high crystallinity and a defect-free crystal. Particle size and shape can be more effectively controlled by preparing MAO-FHA coating through hydrothermal synthesis, and FHA coating with higher crystallinity and uniformity can be 4 ACS Paragon Plus Environment

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obtained on the surface of MAO coating46-49. There are few studies on the preparation of FHA coating on MAO surfaces by hydrothermal synthesis and the degradation mechanism of the composite coating in vitro and in vivo, which is important to evaluate the biocompatibility of the coating. In this study, MAO-FHA coating was prepared on the surface of Mg-3Zn-0.5Zr-0.5Sr alloy using MAO and chemical deposition of FHA. The microstructure, biocompatibility, and corrosion resistance of MAO-FHA composite layers are investigated by in vitro and in vivo experiments. Materials and experimental method Material and sample pretreatment Pure Mg (99.99 wt %) ingot, Zn (99.99 wt.%) particles, and Mg-30 wt.% Zr and Mg-25% Sr master alloys were used as raw materials to prepare Mg-3Zn-0.5Zr-0.5Sr alloy. The raw materials were melted in a vacuum induction furnace (ZG- 10) with mechanical agitation under argon at 720˚C to obtain a cylinder ingot casting. The ingot of Φ60 was homogenized at 300˚C for 24 h and then extruded into a Φ 8 rod using a YQ 32-315 extruder at 380˚C . The bar wire was cut into a sample of Φ 8 × 3 mm. Prior to MAO treatment, the samples were finely polished with 320 #, 800 #, 1500 # SiC sandpaper. The samples were washed with distilled water and acetone to degrease and stored in a dry box. Preparation of MAO-FHA coating The preparation process of MAO-FHA coating is shown in Figure 1. The sample was subjected to MAO treatment using a 50 kW pulse power source (Chengdu Tongchuang Electric Equipment Co., Ltd.). In the MAO process, the Mg-Zn-Zr-Sr alloy sample was used as the anode and the stainless steel plate was used as the cathode inside an electrolyte solution composed of 12 g/L Na3PO4 and 6 g/L NaOH. All samples were treated at a constant voltage of 400 V for 15 5 ACS Paragon Plus Environment


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min with the frequency and duty cycle fixed at 600 Hz and 30%, respectively. The MAO coated samples were removed from the electrolyte solution, cleaned with distilled water and ethanol, and dried at room temperature. Fluorine hydroxyapatite (FHA) coating was prepared on the MAO film layer by chemical deposition. 40 mL 1.29 mmol/L Ca(NO3)2•4H2O was added drop-wise to 40 mL 0.71 mmol/L Na2HPO4, followed immediately by drop-wise addition of 10 mL 0.24 mmol/L NaF and 10 mL distilled water. The pH of the mixed solution was adjusted to 6.6 using a 1 mol/L HNO3 solution and the resulting solution was used as a calcium phosphating solution containing F-. MAO coated samples were placed in the solution heated to 95˚C for 30 minutes prior to being rinsed with distilled water and placed in a dry box. The MAO samples chemically deposited with FHA were designated as MAO-FHA.

Figure 1. Schematic representation of the preparation process of MAO-FHA coating. (a) Microarc oxidation (MAO) process (b) MAO sample (c) Preparation process of FHA coating (d) MAO-FHA samples. 6 ACS Paragon Plus Environment

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Microstructure characterization The microstructure and thickness, as well as the elemental distribution of the coatings were characterized using a field-emission scanning electron microscope (FE-SEM, JOEL 6700F, Japan) and its energy dispersive x-ray energy spectrometer (EDS). The chemical composition of the coating was characterized by X-ray diffraction (XRD, RigakuD/max/2500PC, Japan) using Cu Ka with a small angle at a scanning speed of 5˚/min and scanning angle of 10-80˚. In vitro studies Coating adhesion strength The coating adhesion strength was measured using an automatic scratch tester (WS-2005). The scratch speed was 10N/min, and the tip radius of the head was 0.2mm. The critical force (Wc) value of the coating when it was removed from the substrate was determined. Each sample was tested three times to determine its repeatability. The adhesion strength (F) of the coating is calculated according to the following equation(1) 34,64-65: F=

(1)

Where: H is the Brinell hardness value of the substrate (kg/mm2), and R is the tip radius of the head (mm). The Brinell hardness value of magnesium alloy matrix was 47.6 kgf /mm 2, and the radius of the stylus was 0.2mm. Immersion test In order to test in vitro degradation of the MAO and MAO-FHA coatings, immersion experiments were carried out in simulated body fluid (SBF) solution (6.5453 g·L-1 NaCl, 0.3676 g·L-1 CaCl2•2H2O, 0.3728 g·L-1 KCl, 0.1 g·L-1 Na2SO4, 2.2683 g·L-1 NaHCO3, 0.2681 g·L-1 Na2HPO4, 0.026 g·L-1 NaH2PO4, 6.057 g·L-1 (CH2OH)3CNH2) at pH = 7.4 and 37 ± 0.5˚C for 1 day, 3 days, 7 days, 15 days, 20 days and 30 days. The immersion ratio was 0.05 cm 2/ml. After 7 ACS Paragon Plus Environment


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immersion in SBF, samples were rinsed with distilled water and stored in a dry box. The morphology and elemental distribution of the samples were tested using a FE-SEM (JOEL 6700F, Japan) and EDS. The samples were then immersed in a chromic acid solution (5g K2Cr2O7 is dissolved in a mixed solution consisting of 10 mL H2O and 90 mL H2SO4) to remove corrosion products and residual coating, ultrasonicly cleaned with absolute ethyl alcohol, and dried at room temperature. Three-dimensional morphology and contour map of the samples after corrosion products were observed using a stereomicroscope (LCM, OLS 4000, Olympus, Japan). The weight loss was then calculated by measuring the mass and the corrosion rate using following equation(2): CRavg=

W0  W1  At

(2)

where CRavg is the average corrosion rate, W0  W1 is the weight loss, A is the surface area exposed to SBF, t is the immersion time, and ρ is the density of alloy (≈1.74 g/cm3). The pH value of the SBF solution was measured using a digital pH meter (STARTER 3100, OHOUS). A minimum of three samples at each time point were tested to confirm reproducibility. Electrochemical test Electrochemical testing was performed in a simulated body fluid (SBF) using a Zennium electrochemical workstation (Germany). Electrochemical test was carried out in a standard threeelectrode system. The sample (working area: 0.503 cm2) was the working electrode, the graphite electrode was the counter electrode, and the saturated calomel electrode (SCE) was the reference electrode. The samples were immersed in simulated body fluid (SBF) for 30 minutes and immediately tested. The potential polarization test was carried out at a scanning rate of 1 mV/s. The Tafel curve was fitted by Tafel extrapolation method14,21 (Zview software 3.1) to obtain 8 ACS Paragon Plus Environment

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corrosion potential (Ecorr) and corrosion current density (Icorr). The polarization resistance (Rp) is inversely proportional to the corrosion current density (Icorr). The equation(3) 32,42 is: (3) The results were obtained from the average of three samples in each group.

Cytotoxicity test NIH-3T3 cells were used to assess the cytotoxicity of the samples. The Leaching transmitter was a H-DMEM medium containing 10% fetal bovine serum (FBS). According to the proportion of material weight/extractive transmitter (0.19 g/mL), the standard extractive solution with 100% concentration was prepared and samples soaked for 24 hours in 37˚C. The extract was diluted to a concentration of 75% using H-DMEM medium containing 10% FBS. A 96-well flat-bottomed cell culture dish was seeded at a density of 1000/100 μL and 100 μL of the cell suspension was inoculated per well. The cells were placed in an incubator containing a volume fraction of 5% carbon dioxide gas for 24 hours and then the well medium was replaced. Fresh cell culture medium was added to the blank control group and the positive control group contained cell culture medium with 5% dimethylsulfoxide (DMSO). The test groups contained 100% and 75% extract concentrations . After incubation for 72 hours, cell viability was assessed. The cytotoxicity of the samples was examined by MTT assay. The morphology of the cells was observed with an optical microscope (Nikon Eclipse TE2000-U inverted microscope) and the cells were photographed using an imaging system. 20 μL 5 g/L 3-(4,5-dimethyl-2-thiazolyl)-2,5diphenyl-2-H-tetrazolium bromide (MTT) solution was added to each well and incubated for 4 hours prior to discarding the liquid in the wells. Finally, 150 μL DMSO was added to each well



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and the absorbance at 570 nm was measured on a microplate reader (BIO-TEK ELX 800). The relative value added rate (RGR) of the unit was calculated according to the following equation(4): RGR =A / A0×100%

(4)

where RGR is relative proliferation rate as a percentage, A is the absorbance of the positive test group, and A0 is the absorbance of the blank control group. In vivo experiment Animal models A rod-shaped implant Φ2 × 6 mm was used for in vivo implantation experiments. A total of 12 healthy adult white rabbits weighing 3 ± 0.2 kg were selected for animal testing in the study. The white rabbits were anesthetized with sodium pentobarbital (35 mg/kg). After anesthesia, the surgical site was depilated and peeled. Three holes with a diameter of 2 mm were drilled on the tibia with a hand drill. The MAO coated implant was implanted into the right tibia (Fig. 2a) and the MAO-FHA coated implant was implanted into the left tibia (Fig. 2b). No damage to the coating was observed during implantation. The rabbits were euthanized after 1, 4, and 12 weeks after which the implant and surrounding tissue were removed and stored in a 10% formalin solution. The experiment was completed in the animal room of the Tianjin Institute of Medical Sciences as per standard protocol and guidelines of the Institutional Animal Ethics Committee (IAEC) of TIMS.


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Figure 2. Images of (a) MAO- and (b) MAO-FHA-coated samples implanted in rabbit tibia. Routine pathological examinations Specimens of different implantation cycles were dehydrated and decalcified, embedded in paraffin, and stained with hemoglobin and eosin (H/E) for routine pathological examination. Images were acquired using a high-resolution digital microscopy imaging system (hmias-2000) to observe the formation of bone tissue around the implant. According to GB/T 16886.62015/ISO 10993-6:2007, the tissue response after implantation was evaluated and the degree of inflammatory cell reaction and the formation of fibrous tissue were scored. The evaluation score of the control sample was subtracted from the evaluation score of each observation period sample,and the differences were compared with the standard (see Table S1) to determine the biocompatibility of the coating and the degree of tissue stimulation. The scores are represented by x/y, where x represents the sum of all checkpoint scores and y represents the number of checkpoints.



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Microstructure of implant surface at different implantation times The implants were removed from the bone after 1, 4, and 12 weeks of implantation and the samples were dehydrated using a gradient concentration (including 50%, 70%, 80%, 90%, and 95%) of anhydrous ethanol. After that, the samples were treated with 100% ethanol twice for 20 min each time, and the samples were dried. The dehydrated sample was embedded in an epoxy resin and finely polished using 1500 # SiC sandpaper. The surface morphology, mineralization microstructure, new tissue regions, cross-sectional morphology, and elemental distribution of implants with different implant durations were analyzed using FE-SEM (Quanta FEG 250, USA) equipped with an EDS.


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Results and discussions Characterization of the MAO-FHA coating

Figure 3. XRD patterns of MAO and MAO-FHA coated samples. Figure 3 shows the XRD patterns of MAO and MAO-FHA coatings on the surface of Mg-ZnZr-Sr alloy. MgO diffraction peaks were observed at 2θ = 36.9˚, 42.9˚, and 62.3˚, and Mg3(PO4)2 diffraction peaks were observed at 2θ=32.08˚, 42.51˚, 57.21˚ and 62.24˚. The diffraction peaks of FHA at 2θ = 25.8˚, 31.8˚, 34.1˚, 40˚ and 57.3˚ indicate the FHA coating fully covered the MAO layer. The intensity of the diffraction peak was enhanced at 2θ = 25.8˚, which is (002) for the corresponding crystal plane, indicating the FHA coating exhibited orientation growth in a direction perpendicular to the crystal plane. This observation was in agreement with previous 13 ACS Paragon Plus Environment


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studies21,43. The standard diffraction peak of FHA was slightly shifted to a higher angular position than the HA standard peak. Since the size of F– is smaller than OH–, the substitution of OH– by F– lead to an increase in the original diffraction angle. According to existing crystallographic data, the crystal structure of FHA is similar to that of HA and belongs to the same space group (P63/m; parameters: a=b=9.462 Å and c=6.849 Å, α=β=90˚, γ=120˚)32. The substitution of F– ions with OH– resulted in a-axis shrinkage to 0.9368 nm and no significant change in c-axis size50. Compared with other methods42-44, the FHA prepared by hydrothermal synthesis has better crystallinity21,45-49. Many studies have suggested that higher crystallinity leads to higher bioactivity57.


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Figure 4. Morphology of the (a) MAO-, (c) MAO-FHA-coated samples; (b, d) EDS analysis at Area 1 of (a) and Area 2 of (c), respectively. Figure 4 shows the morphology and EDS elemental spectrums of MAO and MAO-FHA coated samples. The surface of the MAO sample (Fig. 4a) is porous and the diameter of the pores are approx 1-2 μm. This is characteristic of the formation of molten oxide and bubbles releasing during the micro-arc discharge, which can absorb more corrosive electrolytes and reduce the protection of the MAO layer on the Mg matrix51. According to EDS analysis (Fig. 4b), the MAO coating consists of Mg, O and a few P elements. Combined with XRD results, the MAO coating is composed of MgO and a small amount of Mg3(PO4)2. Phosphorous element is derived from the electrolyte solution and magnesium is derived from the matrix, which indicates the 15 ACS Paragon Plus Environment


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magnesium alloy substrate participated in the MAO process. This interaction is advantageous for forming a strong interface bond. After treatment with the calcium phosphate solution containing F–, a uniform flower-like FHA formed on the surface of the MAO-FHA sample and the product is composed of needle-like crystals growing almost perpendicular to the substrate surface (Fig. 4c). Compared with other literatures42,59-60, these FHA have higher aspect ratio and higher similarity with the shape of collagen fibers, especially the tip. EDS analysis (Fig. 4d) showed the MAO-FHA coating contained O, Ca, P, and F elements with a Ca/P atomic ratio of 1.38, indicating the coating was a calcium-deficient FHA. It has been reported that calcium-deficient apatite is more conducive to improving the anti-solubility of the coating and inducing new bone formation in the body32,34,58. The components of FHA are similar to natural bone minerals, which can effectively enhance the bioactivity and cell biocompatibility of the coating51.


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Figure 5. (a) Cross-sectional morphology of (a) MAO and (b) MAO-FHA coated samples; EDS analysis of (b) MAO and (d) MAO-FHA coated samples. The cross-sectional morphology and EDS element analysis of MAO and MAO-FHA coated samples are shown in Figure 5. No break point between the coating and the substrate can be seen from the cross-sectional morphology indicating the coating has good adhesion. The MAO coating was approx 8-9 μm thick and composed of Mg, O and P. The elements were uniformly distributed in the thickness range of the coating, indicating PO43- ion participates in the formation of the coating and the MgO and Mg3(PO4)2 in the MAO coating grew simultaneously. The FHA (Fig. 5c) and MAO-FHA coatings were approx 6 μm and 13-15 μm thick, respectively. EDS 17 ACS Paragon Plus Environment


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analysis shows the MAO-FHA coating was composed of Mg, O, P, Ca elements (Fig. 5d). The intensity of Ca and P increased in the EDS analysis path and their content decreases from the MAO coating. This indicates that the FHA coating has completely covered the MAO coating and increased the density of the coating. Bonding strength The critical strength (Wc) of the MAO and MAO-FHA coated samples was 15N and 10N, respectively. According to the formula34,64-65, it can be calculated that the average adhesion strength of MAO and MAO-FHA coatings is 27.8MPa and 21.6MPa. The ISO 13779-2:200864,65 standard requires a surgical implant with a bond strength of 15 MPa, and MAO-FHA coating has higher coating adhesion strength, which is satisfactory for practical application. Although the bonding strength of MAO-FHA coating decreased, the density and corrosion resistance of the coating increased.


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Growth Mechanism of FHA on MAO Coating

Figure 6. Schematic of the formation of chemically deposited FHA coating on the surface of the MAO sample. Figure 6 shows a schematic of the formation of chemically deposited FHA coating on the surface of a MAO layer. After MAO treatment, a microporous coating with MgO as the main component is formed on the surface of magnesium alloy (Fig. 6a). In the calcium phosphate solution containing F– at 95˚C, the MgO of the MAO coating reacts with H2O to form Mg(OH)252 and hydroxyl groups are introduced on the surface of the sample to 19 ACS Paragon Plus Environment


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raise the pH of the solution. When heated in a water bath, due to electrostatic adsorption, the Ca2+ ions in the solution are adsorbed near the OH– ions and PO43– ions are adsorbed near the Ca2+ ions. As the reaction time prolongs, Ca2+ and PO43– ion concentration tends to be saturated at the surface of the MAO layer and hydroxyapatite (HA) crystal is non-uniformly nucleated and grows in the pores and pore walls of the MAO film layer (Fig. 6b). The coating solution is supersaturated at high temperatures and HA grows rapidly once formed. When the aggregates are grown up, CaF2 is generated by the reaction with the F– ions in solution (Equation 5) (Fig. 6c). Subsequently, CaF2 is further reacted with HPO42– and H2PO4– ions in the solution to generate FHA (Equation 6). Ca5(PO4)3OH +10F− + (7−n)H+ → 5CaF2 + H2O + (3−n) H2PO4− + n HPO42−

(5)

5CaF2 + (3−n)H2PO4− + nHPO42− + (1−x)H2O → Ca5(PO4)3(OH)1−xFx + (10−x)F− + (7−x−n)H− (0

In Vitro and in Vivo Degradation Behavior and Biocompatibility

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