Performance of hydroxyapatite coatings electrodeposited on micro-arc oxidized magnesium alloys using a static magnetic field

Abstract

Biodegradable magnesium (Mg) and its alloy are some of the most widely used functional materials for osteosynthetic applications due to their rapid degradation properties, and thus they do not require surgical removal. However, the rapid degradation of magnesium alloy can cause a high alloy corrosion rate, which needs to be regulated during the bone healing process. In this study, we coated hydroxyapatite (HA) crystal nanostructures on magnesium alloy by an electrodeposition process in the presence of a static magnetic field to inhibit corrosion. The physical and chemical properties of the HA coatings were characterized using SEM, XRD, EDS, as well as a corrosion test. In addition, the interaction between HA coatings and osteoblast cell regulated cellular behavior was investigated. The result indicated that the corrosion resistance ability of the magnesium alloy coated with HA was significantly improved compared with the uncoated magnesium alloy. An initial corrosion potential of Ca–P composite coating at −0.5 V was achieved, which is almost one third of the potential value of the pure magnesium alloy (AZ91D). The proliferation, adhesion and expression analysis of IGF-1 protein indicated that the HA nanocrystals could enhance the viability of the cells. This work provides insight into the development of the next generation of biocompatible alloys for biomedical applications.

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1. Introduction

Magnesium-based implants are attracting considerable attention in biological and biomedical fields, especially for orthopedic applications because of their excellent biocompatibility and biodegradable properties.¹ Compared with titanium-based implants, which are also currently used clinically, magnesium possesses significantly higher bone-implant contact and superior osseointegration,² which improves the biological fixation and anchorage in host bone. However, the low corrosion resistance of magnesium can lead to the loss of mechanical integrity³ as well as disharmony between the kinetics of degradation and bone healing.⁴ To overcome these intrinsic drawbacks of magnesium, various studies aiming to maintain magnesium strength have been carried out,³ and the interaction between the corrosion of the alloy and the association of bone response has also been investigated intensively.^5,6 Among various methods, magnesium modified by alloying has exhibited the potential advantage for their direct contact with the surrounding bone during the degradation,⁵ which reduces the loss of mechanical properties; moreover, the relevant mechanisms have also been studied.⁷ These hard alloy functional materials are different from traditional soft polymer-based biomaterials^8–13 due to their nanoscale structure, mechanical properties, large surface area and unique porous nature. They also exhibit novel chemical and physical properties, e.g. electrochemical properties, which are not usually seen in organic molecular-based polymer materials. In addition, with chemical surface modification, alloy functional materials can mimic the function of extracellular matrix (ECM)¹⁴ to regulate cellular behavior by controlling the interaction at the cell–substrate interface. In particular, magnesium alloys display light density¹⁵ and high Young's modulus, which is similar to natural bone,^1,16 as well as low health risks and adverse effects on the biological system; therefore, they have gained considerable attention in the bone implant field. For example, Kraus et al. investigated degrading magnesium pin implants in a growing rat skeleton, and it was found that the magnesium alloy displayed good osteoconductive properties by enhancing bone accumulation at the surface of the pin.¹⁷ Dziuba et al. developed the magnesium alloy ZEK100 to improve the stability of resorbable materials for osteosynthesis. It exhibited very high initial stability and excellent biocompatibility in the animal model over a period of 9–12 months.¹⁸ Moreover, the pre-treatment of the surface of magnesium alloy is also a promising strategy for overcoming the drawbacks of high corrosion rate.^19,20 Previous studies have suggested that the magnesium-based metal–inorganic matrix materials would further improve the performance of implants in terms of their bioactive and corrosion behavior compared with the magnesium and titanium alloys.^21,22 The debate on element alloying also indicates that the impurities and secondary phases would generate internal galvanic corrosion and cause adverse effects.²³ In addition, the change in pH caused by the liberation of hydrogen would impact the bone reconstruction.²⁴ Among various types of magnesium alloy surface-functional coating technologies, layer-by-layer deposition exhibits distinguished biocompatibility and inhibits gas generation.²² In particular, the HA/MAO composite coatings have been used as an intriguing treatment for improving corrosion resistance and cellular integration.^25,26 Micro-arc oxidation (MAO), a plasma-electrolytic oxidation technology, could remarkably enhance the corrosion resistance of the magnesium alloy in various conditions by creating a barrier on the surface of the metal,^25–27 and MAO-modified magnesium alloys have been widely used in the biomedical and tissue engineering fields.²⁸ Hydroxyapatite (HA) has been the most widely used bioactive inorganic biomaterial. For example, it has been used for osseointegration and fracture healing in clinic for decades.²⁹ However, its relatively low stiffness limits its application to certain extent. In the structure of the HA/MAO, the porous MAO coatings can serve as adequate support for the implantation of HA particles (micro- or nano-sized) to enhance their mechanical properties for biomedical utilization. Although the mechanism associated with the corrosion behavior and biocompatibility of HA/MAO coatings has been extensively discussed in previous studies, the precise mechanism of control of the phase and topography of the HA coatings are still not been fully studied.²⁰ In addition, the topography and phase of materials can regulate cellular behavior; thus, it is necessary to obtain user-controlled coatings with specific properties.

In our study, AZ91 (magnesium alloy containing Al and Zn elements) was employed as the substrate for supporting the HA/MAO composite coatings. Because of its excellent mechanical properties and corrosion resistance, AZ91 has been widely used as the implant material in academic research and clinical applications.^5,30 Magnesium not coated with HA/MAO was regarded as a control sample, and the functional coatings were fabricated by the MAO/EPD method.³¹ In addition, the topography of the HA coatings was modified by a magnetic field, which has been demonstrated as an effective method for regulating the size of HA particles.³¹ The coating performances, such as topography, corrosion resistance and transformation of phase, were evaluated. The interaction between the biocompatibility and performance of the coatings was also analyzed using MTT and osteoblast staining testing, and the relevant mechanism is discussed.

2. Experimental

2.1 Fabrication of HA/MAO coatings

Die-cast AZ91D magnesium alloy (Al 9.22 wt%, Zn 0.72 wt%, Mn 0.418 wt% and Mg balance) was cut into disc shape with 2 × 2 cm², polished by the SiC papers (#200, #600 grits, #1200 grits, respectively), and sonicated in acetone, then dried in N₂ atmosphere for the subsequent anodization. Magnesium alloy, after treatment, was selected as the working electrode and Pt electrode was selected as the counter electrode. The MAO film was fabricated in 10 g l⁻¹ Na₂SiO₃·9H₂O, 5 g l⁻¹ Na₃PO₄·12H₂O, 2 g l⁻¹ NaOH and 1.5 g l⁻¹ CaF₂ solution under pulse voltage mode with 300–400 V for 20–30 min at 20–30 °C. After the anodization process, the samples were ultrasonically rinsed with acetone and distilled water, and then dried in air for electrodeposition.

Uniform HA coatings were formed by electrochemical deposition on MAO films under a static magnetic field. The electrolyte contained 0.1 mol l⁻¹ calcium nitrate (Ca(NO₃)₂), 0.04 mol l⁻¹ sodium nitrate (NaNO₃) and 0.06 mol l⁻¹ ammonium dihydrogen phosphate (NH₄H₂PO₄); moreover, the MAO film served as the cathode and the platinum electrode served as the counter electrode. The pH of the electrolyte was adjusted to 6 using nitric acid, and the electrochemical deposition was achieved at 85 °C with a current density of 10 mA cm⁻² for 60 min. The Teflon electrochemical bath was placed in a static magnetic field, and the direction of the magnetic field in this study was parallel to the electric field (B = 1 T). The magnesium coated with HA/MAO without using a magnetic field was also used as a control.

2.2 Characterization and tests

Immersion tests were carried out in Hank's solution (8 g l⁻¹ NaCl, 0.4 g l⁻¹ KCl, 0.25 g l⁻¹ NaH₂PO₄·H₂O, 0.35 g l⁻¹ NaHCO₃, 0.06 g l⁻¹ Na₂HPO₄·2H₂O, 0.19 g l⁻¹ MgCl₂, 0.19 g l⁻¹ CaCl₂·2H₂O, 0.06 g l⁻¹ MgSO₄·7H₂O and 1 g l⁻¹ glucose, pH 7.8) at 37 ± 0.5 °C according to ASTM-G31-72. In particular, samples were removed after 7 days of immersion, rinsed with distilled water (Di water) and dried in air. The surface morphology of the composite coatings was characterized using scanning electron microscopy (SEM, S-4800, Hitachi, Japan) at 15 kV. The crystalline structure of the sample was analyzed by X-ray diffraction (XRD, D8211, Huber, Netherlands) with a Cu Kα radiation source, and the chemical compositions of the samples were examined by energy dispersive spectroscopy (EDS, EX, Horbia, Japan). Electrochemical measurements were performed using a three electrode system to determine the corrosion resistance ability of samples. A saturated calomel electrode (SCE) and a platinum sheet were used as the reference and counter electrodes, respectively. Potentiodynamic polarization curves were measured at a scan rate of 1 mV s⁻¹ using an electrochemical workstation (CHI 650C, China). The bond between the Ca–P coatings and the alloy were investigated by the adhesive strength test.³²

2.3 Cell culture

Human osteoblasts (bone-forming cells) were purchased from American Type Culture Collection (CRL-11372). In particular, human osteoblast (HOB) cells were cultured in Dulbecco's modified Eagle's medium (DMEM) containing 10% fetal bovine serum (FBS), 1% penicillin, and 1% streptomycin at 37 °C with 5% CO₂ humidified atmosphere. The culture medium was replaced every other day and passaging the cells was subcultured through trypsinization when they reached 80% confluence. All samples (nine examples for each type of implant) were sterilized by ethylene oxide, and then placed in plastic petri dishes with 24 wells. Human osteoblasts were seeded onto Ti, NT and NT/vancomycin (all substrates were 2 cm × 2 cm) with an initial density of 10⁴ cells per cm².

2.4 Cell proliferation and adhesion analysis

The osteoblast adhesion was investigated after 6, 24, and 48 h and the cell viability was determined using the MTT assay. Human osteoblast cells were cultured onto bare Ti, TNT, TNT/vancomycin substrates at an initial density of 1 × 10⁴ cells per cm². After 6, 12, and 24 h, the samples were washed with PBS buffer and transferred to a new 12-well cell culture plate for analysis. 500 μl of fresh culture medium and 0.1 ml of MTT solution were added to the samples. The cells were incubated at 37 °C for an additional 4 h. Then, the MTT-containing medium was removed and 500 μl of dimethyl sulfoxide was added to each well to dissolve all the generated formazan dyes. The optical density of the solution was measured at a wavelength of 570 nm. Human osteoblast cell viability was calculated by the equation: mean OD/blank control mean OD × 100. All MTT assays were repeated for three independent experiments, each experiment contained at least six replicates.

2.5 Cell viability (AO/EB) analysis

Cell viability was determined by the acridine orange/ethidium bromide (AO/EB) assay. The osteoblasts were cultured onto bare Ti, NT, and NT-V substrates with an initial density of 2 × 10⁴ cells per cm². After 12 and 24 h, the samples were washed with PBS buffer. The living and dead cells were stained with AO/EB at room temperature and observed on a CLSM (confocal laser scanning microscope).

2.6 Immunofluorescence analysis of osteogenesis-related proteins

After 7 days, the cells cultured on nHA/MAO, μHA/MAO and uncoated magnesium alloy were fixed in 4% paraformaldehyde in 1 × PBS buffer for 30 min at room temperature. After washing three times with PBS (5 min for each time), the cells were permeabilized in 1% Triton X-100 in PBS for 5 min. Then, the cells were washed twice with PBS (5 min for each time), followed by incubation with 1% BSA/1 × PBS at room temperature. The cells were incubated for 12 h with the addition of primary antibodies (1 : 200) (Gibco, USA) and anti-IGF-1. After washing three times with 0.1 M PBS (5 min for each time), the sections were incubated with a secondary antibody (1 : 200) (Gibco, USA) for 1 h at 37 °C. The HOB cell was stained with DAPI at room temperature, and then observed on a confocal laser scanning microscope for immunofluorescence analysis.

2.7 Statistical analysis

Statistical analysis was performed using a Student's test using SPSS software. The transition of the number of bacteria between each metal and MTT assay (by OD) was analyzed by a variance (ANOVA) test. The data is expressed as mean ± standard deviation, and statistical significance was considered at P < 0.05.

3. Results and discussion

3.1 Morphology of the HA/MAO coatings

The morphologies of three samples were shown in Fig. 1a, the surface of the magnesium was smoothly, and the scratch caused by the polishment could be obviously observed. The HA/MAO coatings fabricated by electrodeposition with or without the presence of a magnetic field exhibited different topography, as illustrated in Fig. 1b and c. The HA particles in the composite coatings were spherical with a diameter of about 30–50 nm; it is interesting to note that the morphology of the coating dramatically changed, exhibiting a broccoli-like shape with less than 5 μm diameters. The transformation of the topography of HA coatings was consistent with our previous studies, specifically the spherical nanoparticles that were formed in the presence of a magnetic field.³¹ The XRD results indicated formation of passivated MgO and each component was distributed in alloy (Fig. 1d). The characteristic HA peaks could be observed in the XRD data of the HA/MAO coatings with or without the presence of a magnetic field. In addition, the nano-size effect and the amorphous phase lead to the broadening of peaks for the nHA/MAO coatings. For supporting the HA coatings, MAO coatings were introduced to enhance the binding between the HA coating and the alloy substrate. The cross sectional morphologies of the MAO layer were investigated by SEM, and the thickness of the MAO layer was found to be about 8 μm (Fig. 1e). It was observed that the Ca–P ceramic-coated micro-arc oxidized magnesium alloys display high porosity structures, and calderas appeared in the layers. The molten oxide particles formed by the sintering process were randomly distributed and formed large aggregates; pores were formed on the molten oxide and entrapment of gas bubbles in the sintered structure was observed.³³ In addition, the peaks shown in the curves corresponded to the crystal plane of Mg₂Ca, Ca₃SiO₅ and Ca(PO₃)₂ (Fig. 1e), respectively, confirming the existence of the composite phase, in addition to the main MgO phase. The presence of Mg₂Ca, Ca₃SiO₅ and Ca(PO₃)₂ indicated the participation of the Si, Ca and P elements in the electrolyte during the chemical reactions.²⁵ Moreover, the EDS results (as shown in Fig. 1e) further confirm the presence of the abundant elements mentioned above. Therefore, the calcium phosphate ceramic layer could provide the favourable gradient coating to electrodeposit Ca–P biological film for further experiments.

Fig. 1 (a) Smooth magnesium surface; (b) spherical nHA particles deposited on the MAO coatings with the synergy effect of the magnetic field; (c) μHA crystals deposited on the MAO coatings without the synergy effect of the magnetic field; (d) the XRD pattern of AZ91 surface, nHA/MAO coatings and μHA/MAO coatings; (e) the surface and cross-section morphology of the MAO coatings, the XRD and EDS of the MAO coatings.

3.2 Performance of the HA/MAO in SBF

The morphology, phase transformation and the passivation potential of the samples was investigated. In the immersion test, it was found that the coatings showed a significant structure change in three samples after 1 week of immersion. For the nHA/MAO, after immersion in SBF for 1 week, the nHA crystal transformed into flake crystals, and the coatings were tightly and neatly packed, as shown in Fig. 2a. For the μHA/MAO coatings, after 1 week of immersion, the HA crystals accumulated into compact cakes and the coatings were less tightly packed than the nHA coatings (Fig. 2b). However, the magnesium alloy was coated with the flocculence sediments and the fractions could be clearly observed (Fig. 2c). In this study, for nHA/MAO and μHA/MAO, the generation and formation of the HA crystals during the calcification process was considered as a nucleation template in SBF incubation, offering hydroxyl terminals that are known as efficient inducers of apatite nucleation.³⁴ These hydroxyl terminals would enroll Ca²⁺ in the SBF first, followed byCl⁻, H₂PO₄⁻ and CO₃²⁻. The same crystal structure as apatite and the rough porous morphologies of the HA obtained during the calcification may also promote apatite nucleation. Some reports demonstrated that most negatively charged groups strongly induced apatite formation.³⁴ In addition, the pH value of the solution would elevate corresponding with degraded substrate in SBF, and then the Ca²⁺ and phosphate ions (such as PO₄³⁻ and HPO₄²⁻) in the solution would react with each other to form abundant Ca–P salt precipitating on the sample. Moreover, the Ca–P coatings formed on the magnesium alloy might go through different transformations of the HA crystal compared with the HA/MAO coatings. The formation of Ca–P group was the first step during the initial immersion period, and the component ions or groups would be absorbed onto the Ca–P group, and the DCP would be formed subsequently, as shown by XRD and EDS analysis in Fig. 2. Specifically, because the deposition of the Ca–P coatings was modulated by the settlement action, the binding force between the coatings and the substrate was lower than that of the HA/MAO coatings; thus, the surface fraction would be observed.

Fig. 2 (a–c) Morphology of the nHA/MAO coatings, μHA/MAO coatings and magnesium surface immersed for 1 week (insets show enlarged SEM images); (d) XRD pattern of the coatings formed on the AZ91 substrate immersed for 1 week; (e) EDS of the coatings formed on the AZ91 substrate immersed for 1 week (inset shows a table of weight% and atom% of different elements); (f) polarization curves of nHA/MAO coatings, μHA/MAO coatings and AZ91 substrate immersed for 1 week in SBF.

To further investigate the Ca–P coating corrosion resistance ability, a potentiodynamic polarization test was performed. Fig. 2e shows the potentiodynamic polarization curves of the untreated magnesium alloy and the two types of HA/MAO layers immersed in SBF. It is worth noting that the corrosion potential of the two types of Ca–P/MAO layers was about −0.5 V (vs. SCE) and −0.75 V, respectively, more positive than that of the magnesium alloy (−1.37 V (vs. SCE)). The slope of the curve of the untreated magnesium alloy sample rapidly increased at the beginning of the anodic side, and then diffusion-controlled anodic current was observed at the end of the curves due to the fast corrosion rate. This indicates that the untreated magnesium alloy substrate suffered severe erosion in SBF compared with the Ca–P coating. The Ca–P coating has a better performance for corrosion resistance due to the sealing layer effect, and the compact coatings further improve the corrosion resistance, which was confirmed by the performance of the nHA/MAO coatings immersed for 1 week. The different polarization behaviors of the Ca–P coatings and the magnesium alloy can be attributed to their different properties and structures. The MAO layer has relatively high porosity and calderas on the surface, as observed in Fig. 1e. Therefore, by increasing the anodic potential during polarization, the corrosive intermediate (Cl⁻) would be rapidly transferred from the outer porous layer to the inner barrier layer of the MAO, resulting in an increase of polarization current. When the polarization current reaches a certain value, the corrosive intermediate reacts with the MAO layer and produces pits on the layer. For the Ca–P coating species, the transfer of the corrosive intermediate (Cl⁻) was inhibited to a certain extent and the increase of polarization current was suppressed during the polarization process due to the covering and blocking effect of the sealing layer. As a result, the corrosion process of the composite coating was delayed.

3.3 Cytocompatibility of the HA/MAO coatings

Osteoblasts are a widely used cell line model in orthopaedic investigations.³⁵ In particular, osteoblasts have been employed to evaluate various alloy components and their potential application for bone implants. Tiainen et al. investigated the biological response of osteoblast cells to carbon-coated TiO₂. It was observed that the carbon coating could promote ALP activity and osteogenic protein gene expression (PCR), as well as enhance osteoblast proliferation.³⁶ In our work, the adhesion and proliferation analysis of osteoblasts was employed to evaluate the cytocompatibility of the HA/MAO coatings. The MTT assay is a colorimetric assay for evaluating the mitochondrial dehydrogenase activity in the cells. In the presence of NAD(P)H-dependent cellular oxidoreductase enzymes, MTT (a yellow tetrazole) could be reduced to its insoluble formazan, which has a purple color. A solution of dimethyl sulfoxide is added to dissolve the insoluble purple formazan product to form a colored solution. The absorbance of this colored solution can be quantified by measuring it at a certain wavelength to determine the mitochondrial activity, which reflects the total cell activity. After a 24 h culture, the nHA/MAO coatings could significantly enhance the cell adhesion compared with the μHA/MAO and uncoated magnesium alloy. Up to 72 h, the enhancement observed was even more obvious. Correspondingly, the MTT results further indicated that the nanostructure crystal HA coatings would considerably improve the activity of the cells (Fig. 3b). In addition, the AO/EB staining demonstrated that all the samples would inhibit the apoptosis of the cells. In particular, the amount of the osteoblasts cultured on the nHA/MAO coatings was the highest, the amount of the cells cultured on micro HA/MAO coatings were second-highest, and the number of cells cultured on the magnesium was the lowest. In addition, to further investigate the consequences for increased cell adhesion and proliferation on the nHA/MAO coatings, we sought to examine the expression of the proliferation-related protein, insulin-like growth factor-1 (IGF-1), by an immunofluorescence analysis. The IGF-1 was the mitogenic factor, which was used to stimulate osteoblast growth and proliferation, and the high expression of the IGF-1 leads to enhanced osseointegration.³⁷

Fig. 3 Behaviors of the osteoblast (CRL-11372) on different coatings: (a) adhesion of cells on nHA/MAO coatings was increased compared with μHA/MAO coatings and magnesium alloy after being cultured for 24 h (*P < 0.05, compared to μHA/MAO coatings and magnesium alloy); (b) the proliferation of the cells cultured on the nHA/MAO coatings was higher than that of μHA/MAO coatings and magnesium alloy after being cultured for 24 h (*P < 0.05, compared to μHA/MAO coatings and magnesium alloy); (c) fluorescence micrographs of the osteoblast cells cultured for 24 h on nHA/MAO coatings, μHA/MAO coatings and magnesium alloy. Living cells (green) and dead cells (red) were stained with acridine orange/ethidium bromide and were visualized using fluorescence microscopy; (d) staining of IGF-1 of cells cultured on three samples.

The expressions of the IGF-1 (Fig. 3d) indicated that the enhanced viability and activity of the cells could be attributed to the higher expression of growth factor in cells cultured on the nHA/MAO coatings. The HA coatings could offer the osteoblast a close native environment; therefore, the osteoblast exhibited increased proliferation and viability on the HA/MAO coatings compared with the smooth magnesium substrate.³⁸ The HA with nanostructure displayed a significant advantage on the osseointegration compared with the micro scale HA crystal. Due to the nano-size structure, the phase of the nHA was amorphous (the broad peak in Fig. 1b), and the dissolution rate was higher than the micro-HA particles;³⁹ thus, the nHA coatings lead to a high extracellular Ca²⁺ concentration, and the high Ca²⁺ concentration enhances the proliferation and viability of the cells by up-regulating the expression of the proteins related to the bone genes such as IGF-1, osteopontin, osteocalcin and bone sialoprotein.⁴⁰ Although the nHA particles exposed to the cells (especially, monocytes or macrophages) would result in the death of the cells,⁴¹ the mode of interaction between nanoparticles and cells would be responsible for the different fates of cells after treatment by the nHA particles. The cells cultured with the gel of nHA particles (especially, the small spherical HA nanoparticles) would result in the uptake of particles in cells;⁴¹ subsequently, the nHA particles would dissolve under the effect of the lysosome and release Ca²⁺ into the cytoplasm, and the intracellular Ca²⁺ homeostasis would be broken by the increased Ca²⁺ ions, which is the main cause of cytotoxicity.³⁹ In our study, the cells were cultured on the nHA coatings, which showed the large agglomeration of HA nanocrystals; moreover, the nHA particles were not dispersed in the medium, and the state of the nHA agglomeration confined the cytophagy (only an agglomeration size smaller than 500 nm could be swallowed by the cell). Moreover, the concentration of the nHA gel, the preparation method as well as the chemical characteristics of the particles would influence the toxicity of the nHA particles: the suspended gel preparation with a concentration of more than 31 μg ml⁻¹ would be toxic, whereas other preparations were also toxic but only at concentrations higher than 250 μg ml⁻¹ The results of our study indicate that the nHA coatings would promote the proliferation of osteoblasts, which is caused by the increased extracellular Ca²⁺ concentration as well as the regulation of the signaling pathway triggered by Ca²⁺. The small amount of nHA particle uptake by the cells might contribute an invalid effect on the cells behavior. In addition, the geometry and scale of nanocoating also should be responsible for the increased adhesion of cells. Previous reports indicate that the nanotopography and the nanopattern coatings (i.e. nanotubes and spherical nanoparticles) could up-regulate the activity of the integrin and adhesive protein.⁴² Thus, it is reasonable to assume that the significant difference in the viability of cells cultured on different coatings could be attributed to the dimensional effect of nanostructure.

4. Conclusions

A calcium phosphate coating was successfully generated on a micro-arc oxidized magnesium surface using an electrodeposition method. With the synergistic effect of a static magnetic field, a spherical HA crystal with nanostructure was achieved. After immersed in SBF, the morphologies of calcification coating have changed dramatically in the different coatings. The nHA coating exhibited excellent corrosion resistance performance due to sealing layer effect. The proliferation and viability of the cells could be improved when cultured on the nHA/MAO coatings due to the high expression level of the IGF-1. Our work provides insight for developing novel clinic-oriented biocompatible functional materials.

Acknowledgements

We thank the Science Foundation of the China University of Petroleum, Beijing (no. 2462014YJRC011) and the National Natural Science Foundation of China (no. 50872018, 51002027), the scientific foundation of the educational department of Liaoning Province (L2012084) and the Project of the Ministry of education of basic scientific research (N130402001) for the support.

Notes and references

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