In Vitro and in Vivo Degradation Behavior and Biocompatibility

Article Cite This: ACS Biomater. Sci. Eng. 2019, 5, 2858−2876

pubs.acs.org/journal/abseba

In Vitro and in Vivo Degradation Behavior and Biocompatibility Evaluation of Microarc Oxidation-Fluoridated HydroxyapatiteCoated Mg−Zn−Zr−Sr Alloy for Bone Application Yansong Wang,† Xiao Li,† Minfang Chen,*,†,‡,§ Yun Zhao,†,‡,§ Chen You,†,§ Yankun Li,† and 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

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‡

S Supporting Information *

ABSTRACT: Magnesium and its alloys are biodegradable materials with great potential for biomedical development; however, their high rate of degradation in biological environments limits 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 microarc oxidation (MAO) combined with chemical deposition of fluoridated hydroxyapatite (FHA). A dense calcium−phosphorus coating 6 μm thick composed of needle-shaped 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 bodily fluids. KEYWORDS: magnesium, microarc oxidation, FHA, biocompatibility, cytotoxicity, in vivo

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INTRODUCTION Biodegradable materials are greatly favorable for patients with bone injuries because secondary surgeries following orthopedic implantation are avoided.1−3 This reduces economic burden, physical distress, and the possibility of various complications for patient.4−7 Magnesium and its alloys are considered to be biodegradable materials with great development potential8−11 especially for bone repair.12 These biocompatible13−15 materials have densities similar to natural bone16,17 and mechanical properties showing no stress shielding effect.18 However, the degradation rates of magnesium and its alloys in the human body are fast,19,20 which has limited the application of these materials.21,22 Surface treatment methods used to improve the corrosion resistance of magnesium alloys include electrochemical deposition,23 chemical conversion coating,24,25 biodegradable polymer coating,26−28 and microarc oxidation (MAO).29 The preparation of a bioactive functional coating on the surface of magnesium alloy, a major focus of magnesium alloy surface © 2019 American Chemical Society

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 treatments.30 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 fluids.31−35,39 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 the HA coating38 and has good bioactivity.31,41 The FHA coating also Received: April 23, 2019 Accepted: May 17, 2019 Published: May 17, 2019 2858

DOI: 10.1021/acsbiomaterials.9b00564 ACS Biomater. Sci. Eng. 2019, 5, 2858−2876

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ACS Biomaterials Science & Engineering

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, and (d) MAO-FHA samples.

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 adhesion.34,61 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 obtained on the surface of MAO coating.46−49 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, the 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.

provides better apatite-like layer deposition, better cell adhesion, better protein adsorption, and improved alkaline phosphatase activity in cell culture, and the F− ion32 of FHA can prevent the decrease of bone density caused by osteoporosis45,62 and also has an inhibitory effect on bacteria.48,63 Therefore, FHA is a potential material for the functional coating of biodegradable magnesium alloy. In general, electrodeposition is used to prepare the FHA coating on the surface of the magnesium alloy. Bakhsheshi-Rad et al.32 electrodeposited the FHA coating on the surface of the Mg−Ca alloy and produced a dense, uniform FHA coating with a slow degradation rate and good apatite-inducing ability. Compared with the 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 the FHA coating on the surface of the 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 condyle.41 The 1 month in vivo animal test showed enhanced interface bioactivity and more direct contacts around the FHA-coated implant. Razaviet 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.

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MATERIALS AND EXPERIMENTAL METHODS

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 diameter 60 was homogenized at 300 °C for 24 h and then extruded into a diameter 8 rod using a YQ 32-315 extruder at 380 °C. The bar wire was cut into a sample of dimensions 8 mm × 3 mm. Prior to MAO treatment, the samples were finely polished with 320 grit, 800 grit, and 1500 grit SiC sandpapers. The samples were washed with distilled water and acetone to degrease and stored in a drybox. Preparation of MAO-FHA Coating. The preparation process of the 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 2859


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Figure 2. Images of (a) MAO- and (b) MAO-FHA-coated samples implanted in rabbit tibia. 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 cm2/mL. After immersion in SBF, samples were rinsed with distilled water and stored in a drybox. The morphology and elemental distribution of the samples were tested using a FE-SEM (JEOL 6700F, Japan) and EDS. The samples were then immersed in a chromic acid solution (5 g of K2Cr2O7 is dissolved in a mixed solution consisting of 10 mL of H2O and 90 mL of H2SO4) to remove corrosion products and residual coating, ultrasonically 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 eq 2:

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 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. A volume of 40 mL of 1.29 mmol/L Ca(NO3)2·4H2O was added dropwise to 40 mL of 0.71 mmol/L Na2HPO4, followed immediately by dropwise addition of 10 mL of 0.24 mmol/L NaF and 10 mL of 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 min prior to being rinsed with distilled water and placed in a drybox. The MAO samples chemically deposited with FHA were designated as MAO-FHA. 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, JEOL 6700F, Japan) and an 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 Kα 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 10 N/min, and the tip radius of the head was 0.2 mm. 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 eq 1:34,64,65 H F= Ä ÅÅ πR2H − Wc ÑÑÉ1/2 ÅÅ ÑÑ Å Ñ ÅÅÇ Wc ÑÑÖ

CR avg =

(W0 − W1) ρAt

(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 the 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 three-electrode 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 min and immediately tested. The potential polarization test was carried out at a scanning rate of 1 mV/s. The Tafel curve was fitted by the Tafel extrapolation method14,21 (Zview software 3.1) to obtain the corrosion potential (Ecorr) and the corrosion current density (Icorr). The polarization resistance (Rp) is inversely proportional to the corrosion current density (Icorr). Equation 332,42 is

(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 kg/mm2, and the radius of the stylus was 0.2 mm. Immersion Test. In order to test the 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

Rp =

βaβc 2.3(βa + βc)Icoor

(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 HDMEM medium containing 10% fetal bovine serum (FBS). According to the proportion of material weight/extractive transmitter (0.19 g/ 2860


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ACS Biomaterials Science & Engineering mL), the standard extractive solution with 100% concentration was prepared and the samples soaked for 24 h 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 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 h, 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% dimethyl sulfoxide (DMSO). The test groups contained 100% and 75% extract concentrations. After incubation for 72 h, 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. In total, 20 μL of 5 g/L 3-(4,5-dimethyl-2thiazolyl)-2,5-diphenyl-2-H-tetrazolium bromide (MTT) solution was added to each well and incubated for 4 h prior to discarding the liquid in the wells. Finally, 150 μL of DMSO was added to each well and the absorbance at 570 nm was measured on a microplate reader (BIOTEK ELX 800). The relative value added rate (RGR) of the unit was calculated according to eq 4: RGR = A /A 0 × 100%

implant durations were analyzed using FE-SEM (Quanta FEG 250) equipped with an EDS.

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RESULTS AND DISCUSSION Characterization of the MAO-FHA Coating. Figure 3 shows the XRD patterns of MAO and MAO-FHA coatings on

(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 2 mm × 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 (Figure 2a), and the MAO-FHA-coated implant was implanted into the left tibia (Figure 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 the guidelines of the Institutional Animal Ethics Committee (IAEC) of TIMS. 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 highresolution digital microscopy imaging system (hmias-2000) to observe the formation of bone tissue around the implant. According to GB/T 16886.6-2015/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. 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 grit SiC sandpaper. The surface morphology, mineralization microstructure, new tissue regions, cross-sectional morphology, and elemental distribution of implants with different

Figure 3. XRD patterns of MAO and MAO-FHA-coated samples.

the surface of the Mg−Zn−Zr−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 studies.21,43 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 the c-axis size.50 Compared with other methods,42−44 the FHA prepared by hydrothermal synthesis has better crystallinity.21,45−49 Many studies have suggested that higher crystallinity leads to higher bioactivity.57 Figure 4 shows the morphology and EDS elemental spectra of MAO- and MAO-FHA-coated samples. The surface of the MAO sample (Figure 4a) is porous, and the diameters of the pores are approximately 1−2 μm. This is characteristic of the formation of molten oxide and bubbles releasing during the microarc discharge, which can absorb more corrosive electrolytes and reduce the protection of the MAO layer on the Mg matrix.51 According to EDS analysis (Figure 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. Elemental phosphorus is derived from the electrolyte solution, and magnesium is derived from the matrix, which indicates the magnesium alloy 2861


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Figure 4. Morphology of the (a) MAO- and (c) MAO-FHA-coated samples and (b, d) EDS analysis at Area 1 of part a and Area 2 of part c, respectively.

μm thick, respectively. EDS analysis shows the MAO-FHA coating was composed of Mg, O, P, Ca elements (Figure 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 15 N and 10 N, respectively. According to the formula,34,64,65 it can be calculated that the average adhesion strength of MAO and MAO-FHA coatings is 27.8 and 21.6 MPa. The ISO 137792:2008 standard64,65 requires a surgical implant with a bond strength of 15 MPa, and the MAO-FHA coating has a higher coating adhesion strength, which is satisfactory for practical applications. Although the bonding strength of MAO-FHA coating decreased, the density and corrosion resistance of the coating increased. Growth Mechanism of FHA on MAO Coating. 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 the magnesium alloy (Figure 6a). In the calcium phosphate solution containing F− at 95 °C, the MgO of the MAO coating reacts with H2O to form Mg(OH)2,52 and hydroxyl groups are introduced on the surface of the sample to raise the pH of the solution. When heated in a water bath, due to electrostatic adsorption, the Ca2+

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 flowerlike FHA formed on the surface of the MAOFHA sample and the product is composed of needlelike crystals growing almost perpendicular to the substrate surface (Figure 4c). Compared with other literature values,42,59,60 these FHA have a higher aspect ratio and higher similarity with the shape of collagen fibers, especially the tip. EDS analysis (Figure 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 antisolubility of the coating and inducing new bone formation in the body.32,34,58 The components of FHA are similar to natural bone minerals, which can effectively enhance the bioactivity and cell biocompatibility of the coating.51 The cross-sectional morphology and EDS element analysis of MAO- and MAO-FHA-coated samples are shown in Figure 5. No breaking 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 approximately 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 (Figure 5c) and MAO-FHA coatings were approximately 6 μm and 13−15 2862


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Figure 5. (a) Cross-sectional morphology of (a) MAO- and (b) MAO-FHA-coated samples and EDS analysis of (b) MAO- and (d) MAO-FHAcoated samples.

Figure 6. Schematic of the formation of the chemically deposited FHA coating on the surface of the MAO sample.

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ACS Biomaterials Science & Engineering 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 concentrations tend to be saturated at the surface of the MAO layer and hydroxyapatite (HA) crystal is nonuniformly nucleated and grows in the pores and pore walls of the MAO film layer (Figure 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 (eq 5, Figure 6c). Subsequently, CaF2 is further reacted with HPO42− and H2PO4− ions in the solution to generate FHA (eq 6).

and MAO-FHA-coated samples. The electrochemical analysis results obtained from the polarization curves are shown in Table 1. The Mg alloy substrate has a low corrosion potential (−1507 mV SCE) and a high corrosion current density (84.9 μA/cm2). After microarc oxidation treatment, the corrosion potential is shifted to the right by about 257 mV SCE, and the corrosion current density is reduced to 1.45 μA/cm2. High corrosion potential and low corrosion current density represent good corrosion resistance, and microarc oxidation treatment greatly improves the corrosion resistance of the magnesium matrix. The reason is that the MAO coating effectively prevents the penetration of the corrosive medium into the matrix and improves the corrosion resistance of the magnesium alloy. Compared with the MAO coating, the MAO-FHA coating has better corrosion resistance. The MAOFHA coating corrosion current density is further reduced to 0.95 μA/cm2. The surface of the MAO film has a porous structure. In the simulated bodily fluid environment, the corrosive medium is easy to enter and causes damage to the coating. After the FHA coating is prepared by hydrothermal synthesis, the surface of the MAO coating becomes dense, further improving the corrosion resistance of the magnesium alloy.37,42 Bakhsheshi-Rad et al.32 also proved that the dense FHA coating can effectively improve the corrosion resistance. In Vitro Degradation Behavior of the MAO-FHA Coating. Figure 8 shows the pH values of the Mg−Zn−Zr−

Ca5(PO4 )3 OH + 10F− + (7 − n)H+ → 5CaF2 + H 2O + (3 − n)H 2PO4 − + nHPO4 2 −

(5)

5CaF2 + (3 − n)H 2PO4 − + nHPO4 2 − + (1 − x)H 2O → Ca5(PO4 )3 (OH)1 − x Fx + (10 − x)F− + (7 − x − n) H−

(0 < x < 1, 0 < n < 3)

(6)

The initially nanosized HA crystals form agglomerated particles, and many FHA crystals are grown in orientation with HA particles and form flower patterns. Meanwhile, the Mg2+ ions near the surface of the microarc oxidation film can also diffuse into the hydroxyapatite lattice and replace the Ca2+ ions in the lattice. Therefore, the calcium/phosphorus atom ratio of FHA formed during the water bath heating process is lower than the standard calcium phosphorus atom ratio of HA (1.67). As the reaction continues, the HA crystal nucleus grows and F− continuously enters the HA lattice and regulated its growth (Figure 6d). To the best of our knowledge, the growth mechanism of FHA coatings prepared by chemical deposition on MAO coatings has not been fully explained in previous studies. Electrochemical Polarization Testing. Figure 7 shows the electrochemical polarization curves of Mg alloy and MAO-

Figure 8. pH of Mg−Zn−Zr−Sr alloy and MAO- and MAO-FHAcoated samples immersed in SBF at 37 °C over time (mean ± SD, n = 3).

Sr alloy and MAO- and MAO-FHA coated samples immersed in the SBF at 37 °C. The magnesium alloy without surface treatment substantially degrades, and the alkalinity of the solution increased to nearly 8.7 after 4 days and 9.35 ± 0.03 after 15 days. The magnesium alloy sample with MAO coating showed comparatively rapid degradation at the beginning of the immersion experiment, and the SBF pH increased to 8.2 ±

Figure 7. Electrochemical polarization curves of Mg alloy and MAOand MAO-FHA-coated samples

Table 1. Electrochemical Analysis Results Obtained from Polarization Curves samples Mg alloy MAO MAOFHA

corrosion potential, Ecorr (mV) vs SCE

current density, Icorr (μA/cm2)

cathodic slope, βc (mV/decade) vs SCE

anodic slope, βa (mV/decade) vs SCE

polarization resistance, Rp (kΩ cm2)

−1507 −1250 −1240

84.9 1.45 0.95

210 262 355

141 35 46

2.20 12.11 22.83

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Figure 9. Weight loss (a) and average corrosion rate (b) of magnesium alloy and MAO- and MAO-FHA-coated samples after SBF immersions (mean ± SD, n = 3).

protective performance than MAO and was shown to effectively improve the corrosion resistance of the alloy substrate. Figure 10 visually compares the three-dimensional morphology of the MAO- and MAO-FHA-coated samples immersed in SBF after removing the corrosion product. It can be seen that the corrosion behavior of the two samples is similar, and the surface of the sample presents uniform corrosion characteristics without significant fluctuation, indicating both MAO and MAO-FHA coatings effectively overcome the pitting corrosion observed in magnesium alloy immersed in SBF. As the immersion time prolonged, the diameter and the integrity of the MAO- and MAO-FHA-coated samples gradually decreased. According to the corresponding contour map, from 3 to 15 days, the diameter of MAO-coated sample decreased from 0.77 mm to 0.7 mm and the diameter of the MAO-FHAcoated sample decreased from 0.78 mm to 0.74 mm. After 3 days of immersion, significant corrosion occurred on the edge of the MAO-coated sample, which became more serious with the extension of time and gradually expanded to the entire sample (Figure 10a−c). This is due to the tip effect in the microarc oxidation process in which the continuous concentration of the electric spark on the edge of the sample results in large holes and cracks allowing the SBF to easily enter the matrix through these defects. The outline of the MAO coating sample immersed in the SBF for 15 days (Figure 10f) shows an obvious left-high, right-low phenomenon indicating the corrosion process is uneven and the MAO coating edge defect was the breakthrough point. Comparatively, the MAO-FHAcoated sample remained relatively complete throughout the immersion process (Figure 10g−i). It can also be seen from the corresponding contour map (Figure 10j−l) that the surface of the MAO-FHA sample was uniform and there was no obvious rugged phenomenon. Therefore, the FHA coating increased the protection of the MAO film layer on the magnesium alloy. The morphologies of the MAO- and MAO-FHA-coated samples immersed in 37 °C SBF for 1, 3, 7, 15, and 20 days are shown in Figure 11. After immersion for 1 day, the MAO sample maintained its original microporous structure; however, the number of pores increases significantly and the size of the pores increases from 1 to 2 μm to 5−6 μm (Figure 11a). This indicates that the edge of the hole in the MAO coating may be the weakest spot and is most likely to react with the SBF. From

0.02 after 4 days. In contrast, the pH of SBF with immersed MAO-FHA samples increased very slowly throughout the experiment. After 4 and 15 days, the pH values were to 7.7 ± 0.01 and 8.1 ± 0.01, respectively. This reveals that the surface pores in the MAO layer of MAO-FHA were covered and sealed by the flowerlike FHA coating, which also presents high stability in the aqueous solution. During the early stage of immersion, this FHA coating effectively prevented SBF infiltration into the matrix. Without this protecting coating, the porous structure of the MAO layer is destroyed by the corrosive medium, which reacts with the surface and the interface between the MAO layer and the substrate and results in defects to form Mg(OH)2. Although the gradual deposition of the Mg(OH)2 corrosion product forms a protective layer on the surface of the MAO, the molar volume of Mg(OH)2 is larger than that of MgO. The Mg(OH)2 deposited on the surface of MAO coating and the presence of Cl− ions in the solution may cause cracks, especially at the interface,53 which may accelerate corrosion. The weight loss and the average corrosion rate of Mg−Zn− Zr−Sr alloy and MAO- and MAO-FHA-coated samples during in vitro testing are shown in Figure 9. It can be seen that the weight loss of uncoated magnesium alloy is much higher than the MAO- and MAO-FHA-coated samples. Due to the formation of metal hydroxide and the increase of the pH value of the solution, the weight loss of the Mg alloy increases sharply.26,27 After 20 days of immersion, the weight loss of the Mg−Zn−Zr−Sr alloy reached 0.095 ± 0.005 g. The weight loss of the MAO coating was higher than the MAO-FHA coating after immersion for 1, 3, 7, 15, and 20 days. After 20 days immersion, the weight loss of the MAO sample was 0.031 ± 0.002 g and the weight loss of the MAO-FHA sample was 0.018 ± 0.004 g. Compared with magnesium alloys, the MAOFHA coated samples showed a significant improvement in corrosion resistance throughout the immersion process (Figure 9b). On the first day, the corrosion rate of uncoated magnesium alloy was 22.06 ± 1.35 mm/year, 4 times that of the MAO coating (5.31 ± 1.26 mm/year) and 6 times that of the MAO-FHA coating (3.69 ± 0.84 mm/year). On the 20th day, the uncoated magnesium alloy corrosion rate was 5.67 ± 0.34 mm/year, 3 times that of the MAO coating (1.86 ± 0.16 mm/year) and 5 times that of MAO-FHA coating (1.10 ± 0.29 mm/year). Due to the protective effect of the uniform and dense MAO-FHA coating, the MAO-FHA coating had better 2865


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Figure 10. continued

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Figure 10. Three-dimensional morphologies (a−f) and contour maps (g−l) of MAO- (a, c, e, g−i) and MAO-FHA- (b, d, f, j−l) coated samples after being immersed in SBF for 3 days (a, b, g, j), 7 days (c, d, h, k), and 15 days (e, f, i, l).

thus these nanoparticles were identified as Ca−P compounds. With the extension of the immersion time, the Ca−P particles were continuously deposited on the surface of the MAO-FHA samples. On the seventh day, the surface of the sample presented a similar flowerlike structure of FHA, and the Ca/P atomic ratio increased to 1.368, which was determined to be calcium-deficient HA. No F element was observed in the EDS analysis of the surface, indicating the layer is grown from the SBF. By comparing the morphology of the Ca−P deposit on the surface of the MAO sample immersed in SBF for 20 days, it can be determined that the MAO-FHA sample has higher surface activity and can rapidly induce Ca−P deposition in SBF. As a result, it is observed that the MAO-FHA surface had no significant changes after immersion for 15 and 20 days compared to that of 7 days, while the HA crystals grew more fully and the protective layer was more complete and dense. With the extension of immersion time, the Ca/P atomic ratio on the surface of the sample gradually increased. Compared with other studies,23,42,58 this study systematically and in detail explained the entire process of the MAO-FHA coating change with SBF immersion. In Vitro Corrosion Degradation Mechanism of MAO and MAO-FHA Coatings. The corrosion degradation process of MAO- and MAO-FHA-coated samples immersed in SBF is shown in Figure 12. The MAO coating is composed of an internal compact layer and external porous layer. During the MAO-coated sample immersion in SBF, the coating degradation and Ca−P composite deposition of SBF solution can occur on the surface and in the holes of the MAO-coated sample. Initially the MAO coating reacts with water in the SBF to form Mg(OH)2 (eq 7) and then reacts with Cl− ions in the solution to form soluble MgCl23,55 (eq 8), which continuously degrades the MAO coating. Mg(OH)2 has the ability to induce Ca2+ and PO43− ion deposition in the SBF solution, enabling Ca−P to form on the surface of the MAO-coated sample and thus has an inhibitory effect on the infiltration of SBF. In theory, the coating degradation and Ca−P deposition occur almost simultaneously, and the dissolution of Mg(OH)2 is slower due to the limitation of Cl− ion concentration in solution.3,55,56 The deposition of fine Ca−P nanoparticles on the surface of Mg(OH)2 has a protective effect, and the coexistence of flaky Mg(OH)2 and Ca−P nanoparticles on the surface of the sample is observed after 1−3 days of immersion. According to experimental results, the corrosion degradation of the MAO-coated samples is uneven, showing severe corrosion

the high-magnification SEM image (Figure 11a, inset), the surface of the coating was covered with fine, uniform particles approximately 40−50 nm in diameter. EDS analysis confirmed these nanoparticles were Ca−P compounds (Table 2). Three days later, large cracks appeared on the surface of the MAO sample and the pores size increased (Figure 11c). The surface of the film layer had many sheets of substances in addition to the Ca−P nanoparticles as shown in Figure 11a. This is a typical morphology of MgO and water reaction product Mg(OH)2 in the MAO film layer and suggests the formation rate of Mg(OH)2 is increasing at this time and grows with the Ca−P compound. After immersion for 7 days, the microporous structure on the surface of the MAO layer disappeared (Figure 11e), the surface is completely covered by clusterlike particles (Figure 11e, inset), and the aggregated material is relatively loose. The EDS analysis (Table 2) shows the surface particles are Ca−P compounds. After immersion for 15 days, the surface morphology of the MAO sample exhibited the growth of Mg(OH)2 encapsulated by Ca−P nanoparticles. After 20 days immersion, a large amount of irregular spherical matter was deposited on the surface of the MAO sample (Figure 11i). At high magnification, numerous nanosized Ca−P compounds are observed (Figure 11i, inset) and no regular voids exist between the spheres. Combined with EDS results, the Ca/P atomic ratio on the surface of the sample can be readily characterized as increasing with the extension of SBF immersion time, indicating the MAO coating provides a good template for the deposition of calcium and phosphorus ions on the surface of the film. The formation of Ca−P compounds thickened over time, thus playing a protective role on the magnesium alloy. Meanwhile, the Ca/P atomic ratio on the surface of MAO coated samples within 15 days of immersion was less than 1, which did not conform to the atomic ratio of the Ca−P compound. It is possible the Mg3(PO4)2 contained in the MAO coating maintains a stable state in the SBF solution, and the relative content of Mg 3(PO 4) 2 increases with the continuous dissolution of MgO on the surface. This would result in a relatively high P content in the EDS analysis. The MAO-FHA sample maintained the original flowerlike morphology of FHA, and the surface roughness increased after immersion for 1 day (Figure 11b). High-magnification illustrations show many nanoparticulate deposits. EDS analysis (Table 2) showed the Ca/P atomic ratio reached 1.206, and 2867


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Figure 11. Morphology of MAO-coated samples (a, c, e, g, i) and MAO-FHA-coated samples (b, d, f, h, j) after immersion in SBF for 1 day (a, b), 3 days (c, d), 7 days (e, f), 15 days (g, h), and 20 days (i, j).

Table 2. EDS Element Analysis of the Box Area in Figure 11 atomic %

atomic %

element

element

sample

O

Mg

P

Ca

sample

O

Mg

P

Ca

Area1 Area3 Area5 Area7 Area9

67.41 64.83 67.17 67.81 59.92

12.74 8.73 11.83 5.48 10.74

14.90 16.51 12.78 13.38 12.70

4.94 9.94 8.22 13.33 16.64

Area2 Area4 Area6 Area8 Area10

68.55 66.40 65.54 50.51 53.60

1.38 0.82 0.91

13.63 15.10 14.59 21.64 19.06

16.44 17.68 19.96 27.85 27.34

on one side caused by edge holes and defects in the MAO coating. When immersed in SBF, the marginal pores and

defects are apt to be reacted. The corrosive medium easily penetrates the film layer through the holes and defects, reaches 2868


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Figure 12. Schematic diagram of the corrosion degradation process of (a−c) MAO- and (d−f) MAO-FHA-coated samples immersed in SBF.

During the immersing process, water and Cl− ions need to penetrate the FHA layer in order to reach the MAO coating (Figure 12d). Due to electrostatic adsorption, Ca2+ ions adsorb in the vicinity of OH− ions on the coating surface at early immersion times, and PO43− and HPO42− ions adsorbed around Ca2+ ions to form Ca−P compounds such as Ca3(PO4)2 and CaHPO4 (eqs 9 and 10) become protective precipitation, which are deposited on the surface of the sample to provide good protection for the MAO-FHA-coated magnesium alloy. When the FHA coating has been corroded by SBF, corrosion products such as Mg(OH)2 are formed under the same mechanism as on the MAO sample and the amount of Mg(OH)2 gradually increases with immersion time. As the volume expansion and internal stress increases, cracks appear in the MAO layer and additional OH− ions are released into the solution, which increases the pH value in the solution (Figure 12e). At the same time, the FHA coating partially dissolves and releases Ca2+, PO43+, F−, etc., which react with OH− (eqs 11 and 12) and form a HA mineralized precipitation layer (Figure 12f).

the interface between the porous layer and the dense inner layer, and causes corrosion degradation inside the layer (Figure 12a). Once Mg(OH)2 is dissolved at the edge defect, the concentration of Cl− ions in the surrounding the SBF solution decreases causing Cl− ions to diffuse toward the defect, which accelerates Mg(OH)2 dissolution at the pores causing defects and weakens the degradation of the still intact MAO coating. Undissolved Mg(OH)2 attracts Ca2+ and PO43− ions from the SBF solution to deposit on its surface and form a protective layer of Ca−P compound on its surface, wherein the consumed Ca2+ and PO43− ions are replenished from other high concentration areas. Finally, uneven macroscopic corrosion degradation occurs, as shown in Figure 10. With the extension of immersing time, the amount of corrosion products such as Mg(OH)2 gradually increases and the micropores are gradually filled with corrosion products and become protective precipitation, which partially reduces the porosity of the MAO layer and protects the magnesium matrix. However, the molar volume of Mg(OH)2 is larger than that of MgO.53 Therefore, as the immersion time is prolonged, Mg(OH)2 accumulated in the pores of the coating and at the interface between the film and the matrix increases the internal stress and causes cracking or shedding of the film layer (Figure 12b). This results in more corrosion ions diffusing into the interior of the MAO layer and dissolving the Mg(OH)2 film layer, and the protective precipitation is broken and loses its protective effect. MgO + H 2O → Mg(OH)2 ↓

(7)

Mg(OH)2 + 2Cl− → MgCl2 + 2OH−

(8)

3Ca 2 + + 2PO4 3 − → Ca3(PO4 )2 Ca 2 + + HPO4 2 − + H 2O → CaHPO4 + 2H 2O

(9) (10)

10CaHPO4 + 2OOH− → Ca10(PO4 )6 (OH)2 + 4PO4 3 − + 10H+ 10Ca 2 + + 6PO4 3 − + 2OH− → Ca10(PO4 )6 (OH)2

As a result, the protective layer of Mg(OH)2 is destroyed, the dense layer of the MAO layer is gradually corroded (Figure 12c), and the corrosion ions penetrate into the substrate and accelerate the corrosion of the matrix. The MAO-FHA coating is composed of a flowerlike FHA layer on top of a MAO porous layer and a dense internal layer.

(11) (12)

Since HA has a relatively low solubility product constant (1.6 × 10−58), Ca−P mineralized precipitation layers are formed on the surface of the MAO-FHA coating and become thicker as the immersion time is extended. The corrosive ions are predicted to enter the interior of the film through cracks and reach the substrate; however, the relatively thick mineralization 2869


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NIH-3T3 cells after 72 h in 75% and 100% extracts (see Figure S2). This indicates the MAO-FHA coating is more effective in inhibiting the release of excessive Mg2+ ions, which is more conducive to cell adhesion and proliferation. Li et al.40 demonstrated that the FHA coating can improve the biological activity of the interface, the corrosion resistance, and the ability of cell proliferation and differentiation. In Vivo Degradation Behavior of MAO- and MAOFHA-Coated Implants. The morphology of MAO and MAOFHA-coated implants are shown in Figure 14. After 1 and 4 weeks of implantation, the appearance of both implants maintained good integrity and their surfaces were covered by biological tissue without obvious corrosion degradation. MAOFHA-coated implants remained uniform throughout implantation (Figure 14b,d). The MAO coated implants showed obvious corrosion pits after 1 week of implantation, indicating corrosion occurred preferentially at the defects of the MAO coating. No substantial corrosion occurred in the implants up to 4 weeks (Figure 14a,c). After 12 weeks of implantation, the pitting area of the MAO coating increased and the bodily fluids had gradually permeated the coating to reach the magnesium alloy substrate, resulting in overall corrosion and degradation of the implant. The surface of the implant was degenerated and cracked, and the average diameter decreased about 1/3 (Figure 14e). Compared with the MAO-coated implant, the MAOFHA-coated implant showed improved surface integrity; however, areas with severe local corrosion and significant reduction were observed (Figure 14f). This indicates the magnesium alloy corroded and degraded, but the degree of

layer slows the rate of corrosion, improving MAO-FHA coating corrosion resistance. Cytotoxicity. Figure 13 shows the cell viability of NIH-3T3 cells in different concentration extracts of Mg alloy and MAO-

Figure 13. NIH-3T3 cell viability in different concentration extracts of Mg alloy and MAO- and MAO-FHA-coated samples (mean ± SD, n = 3)

and MAO-FHA-coated samples. After 72 h incubation, the cell viability of the cells in 100% and 75% concentration extracts was maintained above 75%. Compared with MAO samples, the cell viability of the MAO-FHA samples increased 10% and were consistent with results obtained in optical micrographs of

Figure 14. Morphology of MAO (a, c, e) and MAO-FHA (b, d, f) coated implants after 1 week (a, b), 4 weeks (c, d), and 12 weeks (e, f) implantation. 2870


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Figure 15. Morphologyof MAO-coated (a, c, e) and MAO-FHA-coated (b, d, f) implants after implanting for 1 week (a, b), 4 weeks (c, d) and 12 weeks (e, f).

depositions of Ca−P and a Ca/P atomic ratio of 0.394. After 4 weeks of implantation, the biological tissue on the surface of the implant increased, the Mg peak in the energy spectrum was significantly reduced, and the Ca/P atomic ratio increased to 1.22 indicating the Ca−P compound layer became more thick. The microporous structure of the MAO coating was faintly visible (Figure 15c), indicating that the ability of the MgObased MAO coating to guide new bone growth was weaker than FHA. After 12 weeks of implantation, the morphologies of the two coated implants were similar (Figure 15e,f). The Ca/P atomic ratio of the MAO implant surface coating was 1.537, similar to 1.67 of HA. The Ca/P atomic ratio on the surface of the MAO-FHA implant reached 3.440, indicating the coating may also contain other calcium-containing compounds, further demonstrating the MAO-FHA-coated implant had strong calcification ability and good osteogenic activity. The results were similar to those of Li.41 Figure 16 shows the cross-sectional morphology and EDS line scan elemental analysis of the MAO- and MAO-FHAcoated implants. As the implantation time extended, the thickness of the implant and tissue bonding area of the two coatings were gradually increased and the biological tissue and implant were closely connected. After 1 week of implantation, the surface thickness of the MAO-coated implant was uneven (Figure 16a); the thickest part was approximately 30 μm and the thinnest part was approximately 20 μm. The MAO coating was 10 μm thick and was present with approximately 10−15 μm of biological tissue covering it (Figure 16b). In contrast, the MAO-FHA-coated implants had relatively thick biological tissue growth on the surface with no obvious holes or cracks in the coating and good interaction with the substrate interface (Figure 16c). The EDS line scan shows obvious MAO and

corrosion is significantly less than that of the MAO coated implant. The surface microstructure topography of MAO- and MAOFHA-coated implants (Figure 15) show the two coating were significantly different. After implantation for 1 week, the surface of the MAO-FHA implants did not show the unique FHA flower cluster structure and more biological tissue growth was observed (Figure 15b). EDS analysis contained a Ca−P layer with a Ca/P atomic ratio of 1.417 (Table 3), indicating Table 3. EDS Element Analysis of the Box Area in Figure 15 atomic % element sample

C

O

Mg

P

Ca

Area1 Area2 Area3 Area4 Area5 Area6

51.40 26.34 63.97 67.01 37.68 67.41

34.61 57.33 27.38 25.25 47.05 27.12

7.26 13.84 2.58 0.78 4.63 1.03

4.90 1.03 2.73 2.39 4.19 1.00

1.83 1.45 3.34 4.57 6.44 3.44

the FHA coating has strong biological activity and can rapidly induce calcium and phosphorus to deposit in bodily fluids. The strong Mg peak in the energy spectrum indicates the Ca−P layer deposited after 1 week of implantation was very thin. EDS elemental analysis of the MAO-FHA-coated implant after 4 weeks of implantation (Table 3) showed a weak Mg peak and larger Ca and P diffraction peaks, suggesting the layer of Ca−P on the surface of the implant was thick. The microporous surface structure of MAO remained intact after 1 week of implantation (Figure 15a), with only a few 2871


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Figure 16. continued

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Figure 16. Cross-sectional morphology and EDS line scan elemental analysis of MAO coated- (a, b; e, f; i, j) and MAO-FHA-coated (c, d; g, h; k, l) implants after implanting for 1 week (a−d), 4 weeks (e−h), and 12 weeks (i−l).

FHA double-layer coatings as well as a new tissue layer approximately 15 μm thick composed of Ca, P, and O elements. This may be a Ca−P layer induced by the FHA coating, indicating the MAO-FHA-coated implant had better biological activity than the MAO coating. After 4 weeks of implantation, a substantial amount of biological tissue adhered to the surface of the MAO-coated implant, but the depth of the local corrosion crack was also expanding toward the surface of the Mg alloy (Figure 16e). The biological tissue layer on the surface of the MAO-FHA-coated implant grew thicker. The coating remained uniform and dense, no obvious holes and corrosion cracks were observed, and the interface with the substrate remained intact (Figure 16g). The Mg and O elements on the surface of the MAO-coated implant are observed in the EDS line scan elemental analysis, and the thickness was approximately 3 μm, half as thick as after 1 week. The MAO-FHA-coated implant showed no significant reduction in the thickness of the MgO layer, and the surface Ca−P deposit layer increased to nearly 40 μm, significantly higher than the MAO-coated implant. After 12 weeks of implantation, the MAO coating disappeared and the MAOFHA coating remained scarcely intact (Figure 16i,k). The EDS line scan elemental analysis shows that Mg-based regions appear in the range of approximately 20 μm between the MAO coating and the magnesium alloy, suggesting the magnesium alloy substrate corrosion was relatively serious. Mg and O elements were detected in the range of about 6−7 μm at the interface between the MAO-FHA coating and the magnesium alloy substrate, indicating the FHA coating reduced the corrosion rate of the MAO coating and Mg alloy in contact with bodily fluids. Based on the above experimental results, the MAO-FHA coating was shown to effectively enhance the corrosion resistance of the magnesium alloy. The presence of Ca and P in the MAO-FHA-coated implants increase the deposition rate of Ca−P compounds on the implant surface, stimulate the activity of host osteoblasts, and improve bone conductivity,54 which is beneficial to the growth of bone tissue on the surface of the implant and improves biocompatibility and coating functionality. Biological Evaluation and Histological Observation. The MAO and MAO-FHA-coated implants were implanted into the tibia of rabbits. Postoperative examination revealed no adverse clinical symptoms such as swelling and pain. Tissue reactions scores around the MAO and MAO-FHA-coated implants are shown in Table 4. No cytomegalovirus, fatty infiltration, traumatic necrosis, or foreign body debris were

Table 4. Scores of Tissue Response of MAO- and MAOFHA-Coated Implants time 1 week sample polymorphonuclear leukocyte lymphocyte plasma cell macrophages cytomegalovirus necrosis neovascularization myelofibrosis fatty infiltration traumatic necrosis foreign body debris check point

MAO

MAOFHA

0/2 1/2 0/2 0/2 0/2 0/2 0/2 4/2 0/2 0/2 0/2 2

4 week

12 week

MAO

MAOFHA

MAO

MAOFHA

1/2

7/3

4/3

0/3

0/3

0/2 0/2 0/2 0/2 0/2 0/2 5/2 0/2 0/2 0/2 2

1/3 0/3 1/3 0/3 0/3 4/3 9/3 0/3 0/3 0/3 3

0/3 0/3 0/3 0/3 0/3 6/3 7/3 0/3 0/3 0/3 3

3/3 0/3 3/3 0/3 3/3 2/3 12/3 0/3 0/3 0/3 3

2/3 0/3 2/3 0/3 1/3 4/3 9/3 0/3 0/3 0/3 3

found around the implant at all implantation times, indicating both the MAO and MAO-FHA coatings had good biosafety. The number of immune cells, such as a polymorphonuclear leukocyte, lymphocyte, plasma cell, and macrophages, around the MAO-FHA-coated implant were less than around the MAO-coated implant after 1 week, 4 weeks, and 12 weeks of implantation. Although the number of polymorphonuclear leukocytes is slightly higher at the initial stage of implantation, the data indicates the MAO-FHA-coated implants were less irritating to the surrounding tissue and the degree of fibrosis was relatively low showing the MAO-FHA-coated implants were more biocompatible with the host. After 1 week of implantation, there was no difference in the number of neovascularization in the tissues surrounding the implants. After 4 and 12 weeks of implantation, the number of new blood vessels around the MAO-FHA-coated implant was greater than that of the MAO-coated implant. Neovascularization can promote the formation of new bone, indicating that the MAO-FHA coating has a better role in promoting bone growth. Histologic images of the tissue surrounding the MAOand MAO-FHA-coated implants are shown in Figure 17. The thickness of the myelofibrosis around the MAO-FHA-coated implant after 1 week of implantation is smaller, and the number of bone trabecular is significantly larger than that of the MAO-coated implant (Figure 17a,b). After 4 weeks of implantation, the fibrous layer around both implants was significantly thinner and the number of new bones formed 2873


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had high biosafety and promoted bone growth. More tissue adhered on the surface of the MAO-FHA-coated implants and the bone tissue growth adhesion was thicker, biocompatibility was higher, and the coating was more functional.

■

ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsbiomaterials.9b00564. Histological scoring system; cell viability morphology in different extracts of MAO coated and MAO-FHA-coated samples; EDS element analysis of MAO-coated samples and MAO-FHA-coated samples after immersion in SBF for different times; EDS analysis of MAO-coated and MAO-FHA-coated implants after implanting for different times; NIH-3T3 cell viability cultured in extracts of Mg alloy and MAO- and MAO-FHA coated samples for different times; and hydrogen evolution from different samples in SBF (PDF)

■

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Fax: + 0086-22-6025845. ORCID

Minfang Chen: 0000-0002-7932-9786 Notes

The authors declare no competing financial interest.

■

Figure 17. Histological images of MAO-coated (a, c, e) and MAOFHA coated (b, d, f) implants after 1 week (a, b), 4 weeks (c, d), and 12 weeks (e, f) of in vivo implantation

ACKNOWLEDGMENTS The authors acknowledge the financial support for this work from the National Natural Science Foundation of China Grant (51871166), the Key Projects of the Joint Foundation of the National Natural Science Foundation of China (Grant U1764254), as well as Major Science and Technology Projects in Tianjin (Grant No. 15ZXQXSY00080).

around the MAO-FHA sample was larger (Figure 17c,d). According to previous reports, FHA will directly contact the host bone and the Ca and P elements of FHA can stimulate the osteoblasts to secrete collagen fibers and the mineralization process thereby improving bone conductivity.36 After 12 weeks of implantation, fibrous tissue and new trabecular bone were rediscovered around the two coated samples (Figure 17e,f), which suggests the protective effect of the coating on the magnesium alloy decreased gradually as the implantation time increased and the corrosion degradation rate of the matrix magnesium alloy gradually increased leading to the continuous growth of bone trabecular around the implant. After 12 weeks, although there was minimal tissue necrosis around both implants, regeneration of host tissue was not affected.

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REFERENCES

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CONCLUSION The following conclusions can be drawn: (1) The MAO-FHA coating was successfully prepared on the surface of the Mg− Zn−Zr−Sr alloy by microarc oxidation (MAO) and chemical deposition of fluoridated hydroxyapatite (FHA). (2) When immersed in SBF, the MAO-FHA samples lowered the pH and had slower weight loss and a more complete macroscopic appearance than MAO. The surface of the MAO-FHA-coated implant remained intact and was more capable of inducing hydroxyapatite in the SBF. (3) Compared with the MAOcoated implants, the cell viability of the MAO-FHA-coated implants improved by inhibiting excessive Mg2+ ion release. (4) Compared with the MAO implants, the amount of new bone formation around the MAO-FHA implant increased and the inflammatory reaction decreased. The MAO-FHA coating 2874


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In Vitro and in Vivo Degradation Behavior and Biocompatibility

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