Evaluation of Microstructural Features and in Vitro Biocompatibility of

Article pubs.acs.org/IECR

Evaluation of Microstructural Features and in Vitro Biocompatibility of Hydrothermally Coated Fluorohydroxyapatite on AZ80 Mg Alloy Si-Han Wang,† Chung-Wei Yang,*,† and Tzer-Min Lee‡ †

Department of Materials Science and Engineering, National Formosa University, No. 64, Wunhua Road, Huwei, Yunlin 63201, Taiwan, ROC ‡ Institute of Oral Medicine, National Cheng Kung University, No. 1, University Road, Tainan 70101, Taiwan, ROC ABSTRACT: Mg-based alloys as biodegradable materials have several advantages. However, the extensive applications of Mg-based alloys are limited mainly by their high corrosion rate and loss in structural integrity in a physiological environment. In order to improve the corrosion resistance and biocompatibility of biomedical magnesium, fluorine (F) ions are doped in hydroxyapatite (HA) to form fluorohydroxyapatite (FHA) surface coatings on the Mg-8.5Al-0.5Zn (AZ80) Mg alloy by the hydrothermal synthesizing process. Experimental evidence confirmed that FHA coatings with nanoscaled needle-like crystals can be uniformly deposited without rupture on the entire surface of AZ80 Mg alloy by the hydrothermal synthesizing process. The hydrothermal deposition coating on AZ80 is composed of a Mg(OH)2 intermediate layer and a FHA top coat. X-ray photoelectron spectroscopy results show that fluorine ions are successfully substituted into the HA crystal structure. Potentiodynamic polarization and immersion tests in the Kokubo’s simulated body fluid (SBF) show that the corrosion resistance of AZ80 is significantly improved and the dissolution rate is decreased with the deposition of hydrothermal FHA coatings. The in vitro cell culture studies, using human osteosarcoma MG63 osteoblast-like cell, demonstrated that significant cell viability and proliferation on the surface of FHA-coated AZ80 Mg alloy after 6 to 48 h cell culture. The results suggest that the hydrothermally synthesized FHA coating is effective to improve the in vitro biocompatibility of the Mg-based alloys. reduced.10 In vivo studies indicated that Mg is a basic element in the growth of new bone tissue, metabolism and an essential element of the enzyme system in the human body.11−13 The corrosion layer of Mg alloys is in direct contact with the surrounding bone during degradation, and high Mg ion concentration can induce bone cell activation when Mg alloys are implanted as bone fixtures.14,15 However, the applications of Mg-based alloys are limited due to their high corrosion rate and the biodegradation before the adequate healing of new tissues.16 The degradation rate of Mg-based alloys should be slowed down to allow the mechanical integrity of the implants to remain intact during bone healing. Alloying is one of the effective methods to improve the corrosion resistance and mechanical properties of pure Mg. Therefore, Mg−Al−Zn alloys, such as extruded AZ80A or die-cast AZ91D, and Mg-rare earth (Mg-RE) alloys (e.g., LAE442: Mg-4 wt %Li-4 wt %Al-2 wt %RE and WE43: Mg-4 wt %Y-3 wt %RE) are the mostly

1. INTRODUCTION In orthopedics applications, commonly used 316L stainless steel (ASTM F-138), cobalt−chromium-based alloys (e.g., Co35Ni-20Cr-10Mo, ASTM F-562; Co-28Cr-6Mo, ASTM F799), and titanium-based alloys (Ti6Al4V-ELI, ASTM F-136) play an essential role as metallic implant materials to assist in the repair and replacement of the damaged human hard tissues.1 However, limitations of these metallic implants are the possible release of metallic ions and wear debris, which are generally toxic and harmful to the human body,2−4 as a result of corrosion and wear processes. In addition, serious stress shielding effects5 resulted from much higher elastic moduli of metallic implants compared with human bone tissues will lead to reduce new bone growth and remodeling. In order to avoid these problems, magnesium (Mg) and its alloys have attracted extensive attention nowadays and they are developed as biodegradable implants in the physiological environment.6−9 Mg-based alloys, which have the lowest density (1.74 g/cm3) of constructional metals, have been studied and identified as potential lightweight implants due to their high specific strength (strength-to-density ratio). It is noted that the elastic modulus of Mg alloys (41−45 GPa) is closer to the cortical bone (3−20 GPa) and much lower than titanium alloys (about 110 GPa),6 and the stress shielding effect can be further © XXXX American Chemical Society

Special Issue: International Conference on Chemical and Biochemical Engineering 2015 Received: December 1, 2015 Revised: January 15, 2016 Accepted: February 23, 2016

A

DOI: 10.1021/acs.iecr.5b04583 Ind. Eng. Chem. Res. XXXX, XXX, XXX−XXX

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Industrial & Engineering Chemistry Research investigated Mg-based alloys for biomedical applications6,9,17 because the addition of Al, Zn, and RE elements significantly enhances the mechanical properties by precipitation strengthening effect.7,18 In vitro and in vivo studies also indicate that Mgbased alloys by adding Al or RE elements have an reduced degradation rate, and they can be used for biodegradable medical implants.6,14,15 However, the application of Mg-RE alloys is limited due to the lack of resource and the high cost of RE elements. Even though the Al ions may have neurotoxicity in the human body; however, the considerations for using commercial AZ-series Mg alloys are based on their lower price and better corrosion resistance with preferred mechanical strength than pure Mg. To improve the biodegradation and biocompatibility of Mgbased alloys, forming a surface protective coating is another effective way and calcium phosphate bioceramics have been widely used as coating materials19−23 because of their excellent bioactivity. With the same chemical and crystallographic structure as the major inorganic constituents of human bones, hydroxyapatite (Ca10(PO4)6(OH)2; HA) is a widely used calcium phosphate bioceramic, which is considered an excellent hard tissue substitute in dentistry and orthopedics due to its bioactivity, biocompatibility, and osteoconductive properties.24,25 For clinical applications, HA is commonly used as a bioactive surface coating for the purpose of improving the biological responses and the adhesion of HA-coated metallic prostheses to the human bone.26 However, pure HA coating generally has long-term stability problems in a physiological environment, which results in the bioresorption and causes implant failures.27 In order to improve the chemical stability and corrosion resistance of pure HA, the substitution of cations or anions, such as Mg2+, Sr2+, and F− ions, within the HA crystal structure are developed and studied.28−31 Since fluorine exists in human bones and teeth as an essential trace element against dissolution, fluoridated HA has been developed as a promising candidate for dentistry and orthopedic applications. In this study, a hydrothermal synthesizing process is performed to deposit uniform fluorine-substituted HA (fluorohydroxyapatite, FHA) coatings on the Mg-8.5Al-0.5Zn (AZ80) Mg alloy. The degradation behaviors and electrochemical properties of the FHA-coated AZ80 Mg alloy are examined via immersion and potentiodynamic polarization tests in the Kokubo’s simulated body fluid (SBF). In addition, the osteoblastic cell responses and cell proliferation on coatings are also evaluated to investigate the biocompatibility of the FHA-coated AZ80 Mg alloy.

Analytical grade of dicalcium phosphate dihydrate (DCPD, CaHPO4·2H2O), calcium hydroxide (Ca(OH)2), and hexafluorophosphoric acid (HPF6) were used as reactants. As for the fabrication of fluorine-substituted HA (fluorohydroxyapatite, FHA) coatings, the substitution content of F− ions instead of OH− groups was determined by the x value in the formula of Ca10(PO4)6(OH)2−xFx. HPF6 with concentrations of 1, 2, and 3 M was added into the DCPD/Ca(OH)2 mixture to obtain a solution with Ca/P molar ratio of 1.67, and these F−substituted conditions were labeled as FHA1, FHA2, and FHA3, respectively. The final mixed solution with a pH of 12 was used as the reagent for hydrothermally synthesizing FHA coatings on the AZ80 Mg alloy. For comparison, pure HA coating was designated as a control. The hydrothermal synthesizing FHA process was performed at 175 °C and held for 2 h in a hermetical autoclave (Parr 4621). The heating temperature was maintained throughout the experiments using a heater attached to the autoclave, and the processing temperature was precisely controlled by a proportionalintegral-derivative (PID) controller (Parr 4842) with ±1 °C. The saturated steam pressure was 0.89 MPa in the autoclave at 175 °C. Figure 1 shows the schematic apparatus of substrates setup in the autoclave during hydrothermal synthesizing process.

2. MATERIALS AND METHODS 2.1. Preparation of Hydrothermal FHA Coatings. The base metal (BM) used in this study was 3 mm-thick as-extruded Mg−Al−Zn sheets with a chemical composition of 8.5 Al, 0.5 Zn, 0.25 Mn, 0.023 Si, and Mg balance (in wt %, named AZ80), which was determined by inductively coupled plasma-atomic emission spectrometry (ICP-AES, PerkinElmer/Spectrum one). These sheets were machined into specimens with dimensions of 30 (l) × 15 (w) × 3 (t) mm3 for the hydrothermal synthesis of surface coatings. Before the hydrothermal coating process, AZ80 Mg substrates were ground with 2000 grit SiC papers, grit-blasted with Al2O3 grit to roughen the surface, and then ultrasonically cleaned with acetone and distilled water. The average surface roughness (Ra) of gritblasted AZ80 Mg substrates was controlled at about 4.3 ± 0.4 μm (n = 10).

Figure 1. Schematic apparatus for hydrothermally synthesized fluorine-substituted HA (FHA) coatings on AZ80 Mg alloy within an autoclave.

2.2. Coating Characterization. The phase and composition of the hydrothermally synthesized FHA was identified by a grazing-incidence X-ray diffractometer (GI-XRD, Bruker AXS, D8A25), using Cu Kα radiation at 40 kV and 40 mA over a 2θ range of 20−60° with a scan speed of 3° (2θ) min−1 (step size, 0.02°). The chemical states of F and F− content were determined by X-ray photoelectron spectroscopy (XPS, PHI 5000 Versa Probe) using a standard Al Kα radiation (1486.7 eV) at a pressure of 1 × 10−10 Torr. The surface and crosssectional morphologies of the hydrothermally deposited FHA coatings were examined by scanning electron microscopy B


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Table 1. Comparison of Inorganic Ion Concentrations (mmol/L) for the Used Kokubo’s Simulated Body Fluid (SBF) and Human Blood Plasma Kokubo’s SBF human blood plasma

Na+

K+

Mg2+

Ca2+

Cl−

HCO3−

HPO42−

SO42−

142.0 142.0

5.0 5.0

1.5 1.5

2.5 2.5

147.8 103.0

4.2 27.0

1.0 1.0

0.5 0.5

transferred to new 24-well plates. Subsequently, 200 μL of the 3-[4,5-dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide (MTT) working solution was added to each well and the 24-well plate was incubated for 4 h at 37 °C and 5% CO2. MTT working solution was prepared by dissolving MTT in PBS at 5 mg/mL. At the end of the incubation period, the medium was removed if MTT working with attached cells. Then the 400 μL dimethyl sulfoxide (DMSO) was added to each specimen. After shaking at a speed of 50 rpm for 30 min, the pipet was moved up and down several times to make sure the converted dye dissolves completely. The absorbance of the converted dye was determined at a fixed emission wavelength of 570 nm in an enzyme-linked immunosorbent assay (ELISA) reader (Tecan, Sunrise). Cell proliferation was evaluated by the optical density (OD) value. The experimental data were reported as the mean ± standard deviation (SD). Statistical significant difference was analyzed by one-way analysis of variance (ANOVA) at an average of triplicates, and the statistical significance was defined as p < 0.05.

(SEM, JEOL/JSM-6360). For observation of the cross-sectional morphologies and the coating thickness, the FHA-coated AZ80 specimens were cut by a low-speed diamond saw. The mounted and polished cross sections were coated with gold and then examined by SEM. The porosity of hydrothermally synthesized FHA coatings were quantitatively calculated by an image analyzer (OPTIMAS 6.0). 2.3. Dissolution Test. The standard simulated body fluid (SBF) with a pH of 7.4 was prepared according to the Kokubo’s recipe in which the ion concentrations were similar to the human blood plasma, as listed in Table 1. The dissolution behavior of hydrothermal FHA-coated AZ80 alloy was examined by soaking in the Kokubo’s SBF at 37 °C for different periods of 2, 4, 8, and 16 days according to the ASTM G31-72. After the immersion tests, the samples were taken out, the concentration of Mg2+ was analyzed with ICP-AES, and the surface morphologies were examined using SEM. 2.4. Potentiodynamic Polarization Test. The in vitro corrosion behaviors of uncoated AZ80 alloys and hydrothermal FHA-coated AZ80 specimens within Kokubo’s SBF (pH = 7.4) at 37 °C were investigated by potentiodynamic polarization tests according to the ASTM G102-89. A three-electrode cell was used for potentiodynamic polarization tests. Rectangular specimens with a surface area of 1 cm2 were used as the working electrode, a platinum sheet was used as the counter electrode, and the reference electrode was a saturated calomel electrode (SCE) for the electrochemical measurements. All experiments were carried out in the SBF at a constant scan rate of 5 mV/s. The anodic and cathodic polarization curves were obtained for each specimen. Corrosion current densities and corrosion potentials were determined from the potentiodynamic polarization curves by Tafel extrapolation methods. 2.5. Cell Culture. Human osteosarcoma MG63 cells (American Type Culture Collection, ATCC) were used to assay the osteoblastic cell responses and cell proliferation on the coating surface. MG63 cells were cultured in Dulbecco’s modified Eagle medium (DMEM) containing 10% fetal bovine serum (FBS) and 1% penicillin-streptomycin and maintained in a humidified 5% CO2/balance atmosphere incubator at 37 °C. Subculturing was performed with the use of a phosphatebuffered saline (PBS) with 0.05% trypsin solution. For the cell assay, hydrothermal FHA-coated AZ80 specimens were sterilized and then placed in 24-well culture plates at a density of 1 × 104 cells mL−1 to observe the cell morphology and cell proliferation. After culturing for 3, 6, and 24 h, cells on the specimens were fixed with 2.5% glutaraldehyde for 1 h at 4 °C followed by a series of dehydration with graded ethanol (30%, 50%, 70%, 95%, and 100%, respectively) for 10 min and critical point drying in CO2. Finally, the specimens were gold coated and the cell morphology were observed using a SEM. The MG63 cell viability on the hydrothermal FHA-coated AZ80 Mg alloy was evaluated by a MTT-assay. Each specimen was pipetted with 100 μL of cell suspension (5 × 103 cells/ well). The culture medium was changed every 2 days during culture. After culturing for 6, 24, and 48 h, the culture medium was removed from each well, and the specimens were

3. RESULTS AND DISCUSSION 3.1. Phase Composition. The chemical states and F− content of various hydrothermally synthesized FHA are analyzed by XPS. The survey XPS spectra in Figure 2a demonstrate that Ca, P, F, and O elements are detected. A representative high-resolution F 1s band of various hydrothermally synthesized HA and FHA is shown in Figure 2b. It can be seen that no F 1s XPS spectrum is detected within the HA. It is noted that the binding energy (BE) of F 1s band located at about 684.3 eV can be attributed to the substitution of F−-ions into the FHA crystal structure (the BE of F 1s band for F−-ions in FHA is at 684.3 ± 0.1 eV, while the peak of F 1s band in CaF2 is at 686.7 eV32,33). On the basis of the quantitative XPS analysis results, the chemical formula were obtained and expressed in the formula of Ca10(PO4)6(OH)2−xFx, where x was 0.33, 0.67, and 1.0 for the FHA1, FHA2, and FHA3 conditions, respectively. Figure 3a is the X-ray diffraction pattern of the as-extruded AZ80 alloy. Compared with the standard powder diffraction of Mg (JCPDS 35-0821), the (0002) basal plane, which displays stronger peak intensity, can be recognized as the typical extrusion preferred orientation of the AZ80 Mg alloy. It is recognized that Mg alloys with preferred orientation of (0002) basal-textured orientation exhibited higher corrosion resistance in the physiological environments.34,35 Relatively weak peaks indicated by the triangular marks in Figure 3a are the γMg17Al12 intermetallic compound (JCPDS 01-1128) commonly observed in Mg−Al alloy. Figure 3b shows the XRD patterns of hydrothermally synthesized HA, FHA1, FHA2, and FHA3 conditions. It is recognized that the diffraction peaks of the hydrothermally synthesized HA condition are assigned to a typical pure apatite phase (JCPDS 09-0432). However, the asreceived HA displays lower crystallinity due to its weak diffraction peaks. Besides, a tiny peak of the impurity αtricalcium phosphate phase (α-TCP, Ca3(PO4)2, JCPDS 9-348) is observed in the hydrothermally synthesized HA, as indicated C


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Figure 2. XPS analysis results of the (a) survey spectra and (b) F 1s high-resolution spectra for the hydrothermally synthesized HA, FHA1, FHA2, and FHA3.

Figure 3. X-ray diffraction patterns of (a) the as-extruded AZ80 Mg alloy and (b) the hydrothermally synthesized HA, FHA1, FHA2, and FHA3.

Table 2. Bragg’s Angle for the (300) Diffraction Peak in the XRD Patterns of the Hydrothermally Synthesized Pure HA, FHA1, FHA2, and FHA3

by the circle mark in Figure 3b. The XRD pattern for less F−substituted content of FHA1 (Ca10(PO4)6(OH)1.67F0.33) shows higher crystallinity with a small amount of α-TCP impurity phase. With increasing the F−-substituted content, the XRD patterns for both FHA2 (Ca10(PO4)6(OH)1.33F0.67) and FHA3 (Ca10(PO4)6(OH)F) represent a well-crystallized phase of fluoridated HA (fluorohydroxyapatite, FHA, JCPDS 15-0876), and no other calcium phosphate impurity phases are observed in FHA2 and FHA3 conditions while increasing the F−substituted content. In addition, the sharpening of the three strongest FHA main peaks (for the (211), (112) and (300) at 2θ = 31.9°, 32.3° and 33.1°, respectively) means that the crystallinity and phase stability of FHA can be effectively improved with the substitution of fluorine into the HA lattice.36 It is worth noting that a slight shift of the characteristic (300) diffraction peak from 32.86° of pure HA to 33.06° of FHA3, as listed in Table 2. The XRD data were also analyzed to calculate the lattice constants of these conditions, and the results are shown in Table 3. It can be seen that the a-axis decreased steadily with increasing the degree of fluoridation, while the caxis almost remained unchanged. Concluded from the analysis results, the peak shift is related to the a-axis contraction of the hexagonal apatite crystal lattice because the hydroxyl groups (OH−) is substituted by the smaller size of F− ions to form the FHA crystal.29,37,38 However, the substitution of F−-ions will not apparently change the c-axis of the hexagonal lattice. Thus, the shift of (300) peak to a higher diffraction 2θ angle can be recognized as the demonstration of the substitution of F−-ions into the HA crystal lattice.

peak position (2θ, deg)

HA

FHA1

FHA2

FHA3

32.86

32.92

33.02

33.06

Table 3. Lattice Constants of the Hydrothermally Synthesized HA, FHA1, FHA2, and FHA3 Calculated from the X-ray Diffraction Patterns of Figure 3b lattice constant (Å) a-axis c-axis

HA

FHA1

FHA2

FHA3

9.433 6.885

9.417 6.883

9.389 6.884

9.378 6.883

3.2. Microstructural Features. Figure 4 shows a comparison of the uncoated (grit-blasted) AZ80 Mg alloy, HA-coated, and FHA-coated AZ80 specimens (FHA1, FHA2, and FHA3). We can see that both of the HA and FHA coatings can be uniformly deposited on AZ80 substrate by the hydrothermal synthesizing process. However, the HA coating shows a rougher surface than the FHA coatings. The average surface roughness (n = 10) of hydrothermally synthesized HA and FHA surface coatings on AZ80 substrates are about 5.9 ± 0.5 μm and 3.6 ± 0.3 μm, respectively. The representative surface and cross-sectional SEM images of the hydrothermally synthesized coatings on AZ80 Mg alloy are shown in Figure 5. It can be seen that the hydrothermally synthesized HA coating exhibits a uniform surface morphology, as shown in Figure 5a. D


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synthesized FHA coatings with microstructures illustrated in Figure 5b will provide a bioactive surface to assist in the cell proliferation after implantation. Referring to the abovementioned XPS, XRD analysis, and the microstructural features of hydrothermal synthesized surface coatings on the Mg substrate, the FHA3 (Ca10(PO4)6(OH)F) coating can be recognized as the optimal condition compared to the others in this study. The biological responses and anticorrosive effect of coatings for the AZ80 alloy will be discussed in following sections. Figure 5c,d shows the cross-sectional morphologies of hydrothermally synthesized coatings on the AZ80 Mg alloy. The calculated porosity of HA and FHA coatings are about 10.3% ± 1.0 and 7.2% ± 0.6, respectively. It can be seen that a relatively uniform coating layer with an average thickness of about 250 μm is deposited on the AZ80 by the hydrothermal process. It is noted that the whole hydrothermal deposition coating is composed of a 200 μm Mg(OH)2 intermediate layer and a 50 μm HA/FHA top coat, as indicated in Figure 5c,d. Since the Mg-based alloys have high electrochemical activity and they are quite susceptible to corrosion in solutions, the formation of lamellar Mg(OH)2 layer can be a result from the hydroxylation reaction of the AZ80 Mg alloy. The pH on the AZ80 alloy surface also increases with the hydroxylation reaction. Then the increasing pH promotes the nucleation of crystalline HA/FHA with formation of the Mg(OH)2 layer on the surface.42 With the proceeding of hydrothermal synthesizing reaction, meanwhile, the deposition of HA/FHA coatings is formed with a simultaneous Mg(OH)2 intermediate layer on the AZ80 substrates. 3.3. Electrochemical Corrosion Resistance of the FHACoated/AZ80. Potentiodynamic polarization curve can provide powerful information on the corrosion behaviors and

Figure 4. Photographs of (a) the grit-blasted AZ80 Mg alloy and hydrothermally synthesized (b) HA, (c) FHA1, (d) FHA2, and (e) FHA3 surface coatings on AZ80 specimens.

Through the semiquantitative analysis of SEM/EDS results, the Ca/P molar ratio of the hydrothermal HA coating is about 1.57. A typical surface morphology of the hydrothermally synthesized FHA coatings, as illustrated in Figure 5b, provides evidence of the differences in microscopic surface features compared to the pure HA coating (Figure 5a). The hydrothermal FHA coating, especially for the FHA2 and FHA3 coatings with higher crystallinity (Figure 3b), displays the aggregate of nanoscaled needle-like shape fine crystals, which also uniformly without apparent defects covers the entire surface of the AZ80 substrate. The atomic ratio of Ca/P of those FHA coatings with a needle-like crystal structure is about 1.40. A bit smaller Ca/P ratio indicates that a calcium deficient HA/FHA coating is deposited on the AZ80. It was reported that Ca-deficient apatite and the degree of fluoridation of x = 0.8−1.1 could be more beneficial to improve the dissolution resistance of coatings and the formation of new bone in vivo.28,39 Relative reports also indicated that the coating surface with nanoscaled crystal morphology similar to bone mineral could effectively enhance its bioactivity and cellular biocompatibility.30,40,41 Therefore, it can be expected that the hydrothermally

Figure 5. SEM surface morphologies of the hydrothermally synthesized (a) HA, (b) FHA coatings, and cross sections of the (c) HA, (d) FHA coatings. E


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the corrosion reaction of a metallic substrate is significantly suppressed by the surface treatment. Referring to the observation of microstructural features, it can be recognized that the formation of densely aggregated needle-like fine crystals for the FHA coatings can provide a barrier to prevent the penetration and direct contact of the SBF solution. As a result, FHA coatings with a higher degree of fluoridation represent a much better corrosion potential and a lower corrosion current density than the HA coating. Therefore, the electrochemical testing results demonstrate that hydrothermally synthesized HA and FHA coatings can effectively enhance the corrosion resistance of the AZ80 Mg alloy with a reduced degradation rate, and the protectiveness of the FHA3 coating is better than other conditions. The higher protectiveness of the FHA3 coating is due to its dense microstructural feature with aggregated needle-like fine crystals. In addition, the polarization resistance of the AZ80 Mg alloy (0.35 kΩ/cm2) is also increased with the improvement of corrosion resistance by the deposition of the hydrothermal HA/FHA surface coatings (Table 4). Also, the polarization resistance of FHA2 and FHA3 coatings is 16.04 and 13.77 kΩ/ cm2, respectively. These values are significantly larger than that of the HA coating (1.26 kΩ/cm2). Since the enhancement of the polarization resistance is related to the formation of a surface protective barrier against the corrosive environment for the substrate, the results also confirm that the hydrothermally synthesized HA/FHA coatings can inhibit the penetration of corrosive SBF to the Mg substrate. Compared with the HA coating, the higher corrosion resistance of FHA coatings is a result of its uniform and dense microstructure, which can further reduce the corrosion rate to a lower value. Therefore, it can be conclude that the corrosion rate of AZ80 Mg alloy is effectively reduced with a surface coating, and the FHA coatings with higher degree of fluoridation and densely aggregated needle-like fine crystals commonly display favorable corrosion behaviors in the SBF. 3.4. Dissolution Behaviors. The purpose of in vitro immersion tests in the SBF is to study the long-term stability of the surface coated Mg alloy within a simulated physiological environment. A comparison of the dissolution behaviors of the uncoated AZ80 Mg alloy, hydrothermally synthesized HA, and FHA3-coated AZ80 specimens is shown in Figure 7. The concentration of Mg2+ ions released from AZ80 Mg substrates into the SBF is detected after immersing in the Kokubo’s SBF

corrosion rate. Figure 6 shows the representative potentiodynamic polarization curves of the uncoated AZ80 Mg alloy and

Figure 6. Polarization curves of the uncoated AZ80 Mg alloy and hydrothermally synthesized pure HA, FHA1, FHA2, and FHA3 coatings on the AZ80.

hydrothermally synthesized HA, FHA1, FHA2, and FHA3 coated specimens in the Kokubo’s SBF. The corrosion potential (Ecorr) and the corrosion current density (Icorr) are calculated from the polarization curves using the Tafel extrapolation method. The polarization resistance (Rp), which is inversely proportional to the corrosion current density, is calculated from eq 1. The parameters βa and βc are the anodic Tafel slope and cathodic Tafel slope of the polarization curves, respectively. The corresponding corrosion rate (Pi) is evaluated according to the relationship Pi = 22.85 × Icorr. These electrochemical parameters are listed in Table 4. Rp =

βa βc 2.3(βa + βc)Icorr

(1)

Table 4. Electrochemical Parameters of the Uncoated AZ80 Mg Alloy, Hydrothermally Synthesized HA, FHA1, FHA2, and FHA3-Coated AZ80 in Kokubo’s SBF

AZ80 HA FHA1 FHA2 FHA3

corrosion potential Ecorr (V vs SCE)

corrosion current density Icorr (μA/cm2)

polarization resistance Rp (kΩ/cm2)

corrosion rate Pi (mm/year)

−1.63 −1.45 −1.35 −0.43 −0.26

183.50 19.64 14.37 2.27 4.54

0.35 1.26 3.09 16.04 13.77

4.19 0.45 0.33 0.05 0.10

From the polarization curves, it is observed that uncoated AZ80 Mg alloy exhibits the most negative corrosion potential (−1.63 V) of all the specimens. The corrosion potential of HA and FHA1 (Ca10(PO4)6(OH)1.67F0.33) coated AZ80 specimens slightly shifts to −1.45 V and −1.35 V, respectively. It is noted that the corrosion potential is significantly improved to −0.43 V for the FHA2 (Ca10(PO4)6(OH)1.33F0.67) and −0.26 V for the FHA3 (Ca 10 (PO4 ) 6(OH)F) specimens. Meanwhile, the corrosion current density is dramatically decreased from 183.5 μA/cm2 for the uncoated AZ80 alloy to much lower values (Table 4) with the deposition of HA/FHA coatings. In general, a shift of the corrosion potential toward the positive side and a decrease of the corrosion current density mean that

Figure 7. Dissolution behavior of the Mg2+ ions for the uncoated AZ80 Mg alloy, HA-coated and FHA3-coated specimens in the Kokubo’s SBF. F


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Industrial & Engineering Chemistry Research for 2−16 days. As a result, the amounts of Mg2+ ions increased with immersion time for all the specimens. It is obvious that Mg undergoes active dissolution below pH 10.5. Thus, the corrosion of AZ80 alloy progresses continuously in the SBF. However, the dissolution of AZ80 alloy is significantly suppressed with the deposition of a surface coating, and the dissolution rate of the FHA3-coated AZ80 is slower than that of HA-coated AZ80 after 4−16 days immersion. The lower dissolution rate of the FHA3 coating, compared the HA coating, can be recognized as its unique microstrucre with densely aggregated needle-like crystal layers, which can provide better protection and long-term stability for the substrate. The dissolution behavior can be correlated with the electrochemical measurements represented in Figure 6 and Table 4. Since the Mg(OH)2 layer is porous with lamellar microcracks, the increases of Mg2+ ions concentration is released from the continuous corrosion of Mg alloy and the dissolution of Mg(OH)2 in the SBF. Relative reports indicated that the fluoridated HA crystal structure displays better chemical stability and even lower solubility than stoichiometric pure HA in a physiological environment.28,30,38 Referring to the SEM cross-sectional images, a HA/FHA top coat is deposited on the Mg(OH)2 layer. Therefore, it is resonable to deduce that a more chemically stable FHA top coat maintains the microstructural integrity to effectively reduce the further penetration of corrosive solution and the release of metallic ions from the substrate. Figure 8 shows the morphologies of hydrothermal FHA3coated AZ80 specimens after immersion in the Kokubo’s SBF

and 24 h of culture on the FHA3-coated AZ80 specimens. As shown in Figure 9b, the cell maintains a spherical shape; however, it attaches on the surface with filamentous pseudopodia extensions, which is denoted by a triangular mark, for 3 h of culture. With increasing the incubation time from 6 to 24 h, the cells are observed spread out and attach well on the FHA3 coating with their lamellipodium, as shown in Figure 9c,d. It is reported that the corrosion rate and release of Mg2+ ions have a significant influence on the cell viability, as well as on cell attachment and spreading.43 In comparison with the cell morphologies for 24 h of culture on HA-coated surface (Figure 9a) and FHA3-coated surface (Figure 9d); therefore, it can be seen the osteoblastic cells appear better extensions and display good viability on the FHA coating. Concluded with the polarization curves and immersion tests, the above results reveal that the effectiveness of incorporating F− ions into the HA crystal structure can not only improve the corrosion and dissolution behaviors but also enhance the response and adhesion of osteoblastic cells. As shown in Figure 10, the cell growths of MG63 on the surface of FHA3-coated AZ80 specimens are evaluated by MTT assay after culturing for 6, 24, and 48 h, respectively. Cells are also cultured in the polystyrene served as the control group. For 6 h of culture, the cell numbers indicated by the higher optical density (OD value) have not shown an evident difference between the polystyrene (PS) control and the FHA3-coated AZ80. After 24 and 48 h of culture, the cell numbers on the FHA3-coated AZ80 are significantly higher (p < 0.05) than those on the PS control, indicating that the fluoridated HA coatings have better biocompatibility. In addition, the apparent increases in cell numbers (p < 0.05) are found for the FHA3-coated AZ80 from 6 to 48 h cell culture. It is recognized that the degradation of Mg-based alloys in physiological environments commonly leads to change in Mg2+ ions concentration and increases the pH value, which result in a negative effect on the cell viability. As a result of potentiodynamic polarization tests (Figure 6 and Table 4), the dissolution rate of AZ80 Mg alloy is significantly suppressed with the deposition of a fluoridated HA coating. Furthermore, the substitution of F− ions within the HA crystal structure can effectively improve the chemical stability and dissolution behaviors of the fluoridated HA.30,44 Therefore, the abovementioned factors supply strong beneficial effects to enhance cells proliferation on the FHA-coated (especially for the FHA3 condition, Ca10(PO4)6(OH)F) AZ80 Mg alloy in this study.

Figure 8. Photographs of the FHA3-coated AZ80 specimens after immersion in the Kokubo’s SBF for (a) 2 days, (b) 4 days, (c) 8 days, and (d) 16 days.

for 2−16 days. It can be seen that the hydrothermal FHA3 coating still possesses structural integrity and uniformly covers the whole surface of AZ80 alloy without evident dissolution and dissociation after immersion for 2−8 days (Figure 8a−c). Even though the FHA3-coated AZ80 was immersed for 16 days, no serious coating rupture with the appearance of AZ80 substrate, which resulted from the corrosion in the SBF, is observed on the FHA3-coated specimen, as shown in Figure 8d. The abovementioned result demonstrates that the FHA3 coating with densely aggregated needle-like fine crystal layers displays favorable corrosion resistance and long-term stability within the physiological environment. 3.5. Cell Responses on the FHA-Coated/AZ80. Figure 9a shows SEM micrographs of MG63 osteoblastic cells after 24 h of culture on the HA-coated AZ80 specimens. It is observed that the cells are spherical in shape and attach onto the surface without evident extension at the initial 24 h. Figure 9b−d shows the representative morphologies of MG63 cells after 3, 6,

4. CONCLUSIONS In this study, the microstructural features, corrosion resistance, and in vitro cell responses of the hydrothermally synthesized fluorohydroxyapatite-coated AZ80 Mg alloy have been characterized. The experimental results can be summarized as follows: (1) Uniform FHA surface coatings can be successfully deposited without decohesion on the AZ80 Mg alloy by the hydrothermal method. (2) XPS analysis demonstrates that fluorine ions are substituted into the HA crystal structure to form Ca10(PO4)6(OH)2−xFx with a fluoridated degree (x value) of 0.33, 0.67, and 1.0. (3) The corrosion resistance of the AZ80 Mg alloy is significantly increased with the deposition of hydrothermal surface coating. The better corrosion resistance and lower degradation rate can be achieved at a higher fluoridate degree of G


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Figure 9. Morphologies of MG63 osteoblastic cells attached on the (a) HA-coated AZ80 alloy after 24 h of culture. The cells attached on FHA3coated (Ca10(PO4)6(OH)F) AZ80 alloy after (b) 3, (c) 6, and (d) 24 h of culture.

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AUTHOR INFORMATION

Corresponding Author

*Phone: +886-5-6315478. Fax: +886-5-6361981. E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This study was financially supported by the Ministry of Science and Technology, Taiwan, ROC (Contract No. MOST 1032221-E-150-004) for which we are grateful.

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Figure 10. MTT optical density of MG63 osteoblastic cell proliferation on the polystyrene (selected as the control) and the FHA3-coated AZ80 Mg alloy after 6, 24, and 48 h cell culture (PS, polystyrene; one-way ANOVA, *, p < 0.05).

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Evaluation of Microstructural Features and in Vitro Biocompatibility of

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