Mussel-inspired nano-multilayered coating on magnesium alloys for enhanced corrosion resistance and antibacterial property

Colloids and Surfaces B: Biointerfaces 157 (2017) 432–439

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Mussel-inspired nano-multilayered coating on magnesium alloys for enhanced corrosion resistance and antibacterial property Bi Wang a , Liang Zhao b , Weiwei Zhu a , Liming Fang b , Fuzeng Ren a,∗ a

Department of Materials Science and Engineering, South University of Science and Technology of China, Shenzhen, Guangdong 518055, China Department of Polymer Science and Engineering, School of Materials Science and Engineering, South China University of Technology, Guangzhou 510641, China b

a r t i c l e

i n f o

Article history: Received 22 March 2017 Received in revised form 2 June 2017 Accepted 15 June 2017 Keyword: Magnesium alloy Polydopamine Carbonated apatite Sliver Antibacterial

a b s t r a c t Magnesium alloys are promising candidates for load-bearing orthopedic implants due to their biodegradability and mechanical resemblance to natural bone tissue. However, the high degradation rate and the risk of implant-associated infections pose grand challenges for their clinical applications. Herein, we developed a nano-multilayered coating strategy through polydopamine and chitosan assisted layer-bylayer assembly of osteoinductive carbonated apatite and antibacterial sliver nanoparticles on the surface of AZ31 magnesium alloys. The fabricated nano-multilayered coating can not only obviously enhance the corrosion resistance but also significantly increase the antibacterial activity and demonstrate better biocompatility of magnesium alloys. © 2017 Elsevier B.V. All rights reserved.

1. Introduction Magnesium (Mg) alloys are promising candidates for loadbearing orthopedic implant materials due to their biodegradability, mechanical compatibility and biological safety which can avoid second removal and reduce the stress-shielding effect [1–3]. However, the rapid degradation of Mg alloy in the physiological environment does not match the healing or growth of new bone tissues which will often cause hydrogen accumulation and local alkalization [4]. Various surface modification strategies have been proposed to increase the corrosion resistance of Mg alloy [5–12]. Among them, alkaline-heat treatment [13] and calcium phosphates coatings [8,14,15] are two typical effective methods. On the other hand, for implant biomaterials, the possibility of bacterial adhesion to the surface of implant and subsequent biofilm formation at the implantation site can lead to implant failure [16,17]. The above mentioned two typical coating strategies usually lack of antibacterial activity. In order to prevent such implant-associated infections, various techniques including thermal spraying [18], electrochemical deposition [19–21] and magnetron co-sputtering [22], have been used to incorporate antimicrobial agents, such as silver, copper and zinc ions, into the

∗ Corresponding author. E-mail address: [email protected] (F. Ren). http://dx.doi.org/10.1016/j.colsurfb.2017.06.013 0927-7765/© 2017 Elsevier B.V. All rights reserved.

surface coating [17,18,23–26]. These coatings have demonstrated good osteoconductivity and enhanced antibacterial activity. Inspired by the composition of adhesive proteins in mussels, dopamine self-polymerization has been found to be able to form a thin, surface-adherent polydopamine (PDA) coating on a variety of material [27,28] and such PDA coating can further enhance the surface bioactivity [29] and promote cell adhesion and viability [30]. PDA-induced nanocomposite Ag/calcium phosphates coatings on the surface of titania nanotubes has demonstrated excellent antibacterial and osteointegration functions. In our previous work, with the aid of PDA, strontium substituted hydroxyapatite nanocrystals and bisphosphonate alendronate have been coimmobilized on the surface of AZ31 Mg alloy for osteoporotic repair [31]. Considering that carbonated apatite (CAp) nanoparticles have superior biocompatibility, bioactivity and osteoconductivity due to chemical and structural similarity to biological apatite [32], and that chitosan has excellent biocompatibility and biodegradability and has been used as antimicrobial agent [33], to simultaneously increase the corrosion resistance and antibacterial activity while enhance good biocompatility of magnesium alloys, herein, we developed a facile but effective nanocomposite coating strategy through PDA and chitosan assisted layer-by-layer assembly of biomimetic carbonated apatite (CAp) and sliver nanoparticles on the surface of AZ31 magnesium alloys, as schematically illustrated in Fig. S1. After alkaline pre-treatment, a PDA transition layer was deposited on the AZ31 Mg alloy substrate. Then, a chitosan intermediate layer was grafted through Michanael addition

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and/or Schiff-base reactions between the amino groups of chitosan and the catechol/quinone groups in PDA, followed by immobilization of a layer of PDA coated CAp (CAp@PDA) nanoparticles, another chitosan intermediate layer and a layer of PDA coated Ag (Ag@PDA) nanoparticles. Such chitosan-CAp@PDA-chitosanAg@PDA periodic unit can be repeated for the required number of cycles. The proposed coating strategy have effectively combined the advantages of the antibacterial activity of silver nanoparticles, the excellent bioactivity and osteoconductivity of biomimetitic CAp nanoparticles, the strong adesion of PDA and the biocompatibility of chitosan. The obtained nano-multilayered coating thus have multifunctions: (1) obviously increase the corrosion resistance of the magnesium alloy substrate; (2) significantly increase the antibacterial activity of the implants; and (3) effectively improve the biocompatility. Such coating strategy overcomes the conflicts between antibacterial property and biocompatibility, provides deep insights into surface functionalization of magnesium alloys and can potentially be extended to design of many other orthopedic devices.

2. Materials and methods 2.1. Materials AZ31 Mg alloy ingot was supplied by National Engineering Research Center for Magnesium Alloys, P. R. China and the asreceived ingot was cut into discs with 10 mm in diameter and 2 mm in thickness using a wire-cut electrical discharge machine. Dopamine hydrochloride, tris(hydroxymethyl) aminomethans (Tris) (99.0%), chitosan (low molecular weight, deacetylation degree = 98%), silver nitrate (AgNO3 , ≥99.8%), NaHCO3 , Na2 HPO4 , Ca(NO3 )2 ·4H2 O, and 3-(4,5-dimethylthiazol-2yl)2,5-diphenyltetrazolium bromide (MTT) were purchased from Sigma–Aldrich (USA). Fetal bovine serum (FBS), a-minimum essential medium (␣-MEM), and 1% penicillin−streptomycin solution were purchased from HyClone (USA) and used without further purification. The water was purified before use in a three stage ® Sartorius arium pro purification system and has a resistivity higher than 18.2 MV cm−1 .

2.2. Alkaline treatment of AZ31 Mg alloy substrates The AZ31 Mg alloy discs were ground with SiC emery paper down to a 1200-grit and then polished using 1 ␮m size alumina suspension. After ultrasonically cleaned in double distilled water, ethanol and acetone for 10 min, respectively, the discs were immersed in 100 mL of 0.5 M NaHCO3 solution for 24 h, rinsed with double distilled water three times and dried in air.

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2.4. Preparation of Ag@PDA nanoparticles Silver nanoparticles were prepared by the reaction between 20 mL AgNO3 (1 mM) aqueous solution and 2 mL (40 mM) sodium citrate aqueous solution at 100 ◦ C for 1 h under vigorous stirring. Silver nanoparticles were formed when the color of the solution changed into yellow-green. Then, the precipitates were collected by centrifugation (8000 rpm, 20 min) and washed with double distilled water three times. The Ag@PDA nanoparticles were prepared following the identical procedure with that of CAp@PDA.

2.5. Assembly of CAp@PDA and Ag@PDA on AZ31 magnesium alloy substrates The alkaline pre-treated magnesium alloy discs were sequentially immersed into the following solutions: (1) 2 mg/mL dopamine solution for 24 h; (2) 4 mg/mL chitosan solution (pH = 5.0) for 15 min; (3) 4 mg/mL CAp@PDA suspension for 15 min; (4) 4 mg/mL chitosan solution (pH = 5.0) for 15 min; and (5) 2 mg/mL Ag@PDA solution. All the treated specimens were cleaned ultrasonically for 2 min with double distilled water before moving to the next step. Such chitosan-CAp@PDA-chitosan-Ag@PDA periodic unit can be repeated for the required number of cycles. The corresponding nano-multilayered coatings are denoted (chitosanCAp@PDA-chitosan-Ag@PDA)n , where n is the number of cycles. Here, we use n = 3. For the antibacterial activity and biocompatibility evaluations, we also switched the order to assemble Ag@PDA and CAp@PDA layers. The coated sample with CAp@PDA as the outmost layer ((chitosan-Ag@PDA-chitosan-CAp@PDA)3 ) was also prepared as control.

2.6. Materials characterization The size and morphology of CAp, CAp@PDA, Ag and Ag@PDA nanoparticles were examined by transmission electron microscopy (TEM) and high resolution TEM (FEI Tecnai F30). The Ag nanoparitcles were confirmed by UV–vis spectroscopy (Lambda 25, PerkinElmer) and their size distrution was measured by dynamic light scattering (DLS, Horiba Scientific) technique. The surface composition was analyzed by X-ray photoelectron spectroscopy (XPS, Kratos Axis Ultra DLD, Japan) using monochromatic Al K␣ radiation. All binding energies were calibrated with reference to the C 1s hydrocarbon peak at 284.6 eV. Surface morphology of the coatings was examined by scanning electron microscopy (SEM, Zeiss Ultra 55). The cross-section of the multilayered coating was prepared by focused ion beam (FIB, FEI Helios 600i). To protect the original surface from Ga ions damage, a platium cap layer with thickness of ∼1 ␮m was first deposited before milling the trench.

2.7. Static immersion in the simulated body fluid (SBF) 2.3. Preparation of CAp@PDA nanoparticles First, biomimetic CAp nanoparticles (with the obtained carbonate content of 3.7 wt.%) were synthesized via a hydrothermal route, as detailed in our previous work [32]. The hydrothermal reaction was conducted at pH = 10 and temperature of 200 ◦ C for 24 h. The obtained CAp nanoparticle were dispersed into 2 mg/mL dopamine in a Tris-HCl buffer (10 mM, pH = 8.5) under ultrasonication for 30 min. The mixture was stirred vigorously for 12 h at 60◦ C, and then subjected to centrifugation (speed at 8000 rpm for 20 min) and finally washed with double distilled water and ethanol three times, respectively.

In vitro degradation and ions release were determined by immersing the mechanically polished AZ31 and coated AZ31 discs with Ag@PDA as the outmost layer into 5 mL of SBF at 37 ◦ C for 14 days. The SBF was refreshed at each pre-defined intervals. The concentrations of Mg2+ and Ag+ ions in SBF were measured by inductively coupled plasma mass spectrometry (ICP-MS, Agilent 7700X). The pH evolution was investigated by immersing the two groups of samples into 110 mL SBF (with sample surface area/volume of solution = 1:50) at 37 ◦ C in a water bath for 14 days. During the immersion period, the pH of the SBF was monitored by a pH meter. For each group, three individual measurements were performed.

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2.8. Electrochemical corrosion tests The effect of coating on the corrosion resistance of AZ31 Mg alloy was evaluated by electrochemical potentiodynamic polarization in a conventional three-electrode system using a Princeton Applied Research Versa Studio (PARSTAT 4000). A platinum plate served as counter electrode and a saturated calomel electrode (SCE) as the reference electrode. The samples were insulated with epoxy resin with 78.5 mm2 exposed area as the working electrode. Measurements were performed at room temperature in SBF. The working electrode was immersed in SBF for 1 h to attain a stable open circuit potential (OCP). Potentiodynamic polarization curves (Tafel plots) were recorded by scanning from−0.25 V to + 0.8 V at a constant scanning rate of 1 mV/s. After completing the polarization tests, the samples were washed in double distilled water and ethanol, and then dried in a vacuum for 1 day prior to further examination by SEM. 2.9. Antibacterial assays E. coli (ATCC 25922) and S. aureus (ATCC 29213) were selected as the model Gram-negative and Gram-positive bacteria, respectively. For the bacterial culture, two bacterial suspensions were prepared by firstly taking a single colony from the stock bacterial culture with a loop, and then inoculating in 5 mL sterile nutrient broth medium. After 12 h incubation in a constant temperature vibrator at 37 ◦ C with a speed of 200 rpm, both bacterial species at the exponential growth phase were harvested. To evaluate the antibacterial activities of the coated AZ31 and the bare AZ31 Mg alloys as control against E. coli and S. aureus. Firstly, 1 mL of the bacteria suspension with a concentration of 1 × 107 colony forming units (CFU/mL) in phosphate buffer saline (PBS) was added into each well in 24-well plated containing samples and incubated for 2 h and 4 h at 37 ◦ C, respectively. Then the planktonic bacteria in culture medium were analyzed by the spread plate method. The solutions were serially diluted 1000-fold, and plated in triplicate onto sheep blood agar (SBA) and incubated at 37 ◦ C for 24 h, then the number of CFUs on the SBA of 10 ␮L was counted. The antibacterial rate for planktonic bacteria (Rp ) was calculated based on the formula Rp % = 100(B−E) . Where E repreB sents the average number of viable bacteria in the culture medium incobuated with the coated AZ31 specimen, B is the average number of viable bacteria of the bare AZ31 as control. After the bacteria were incubated with the three groups of samples (bare AZ31, coated AZ31 with Ag@PDA as the outmost layer, and coated AZ31 with CAp@PDA as the outmost layer) for 2 h and 4 h at 37 ◦ C, respectively. The samples were rinsed three times with PBS and fixed with 2.5% glutaraldehyde in PBS for 1 h at room temperature. The samples were then rinsed antoher three times with PBS, dehydrated in graded ethanol (30%, 50%, 75%, 95% and 100% ethanol for 10 min, respectively), critical point dried, gold sputtered and observed with SEM. 2.10. Osteoblastic MC3T3-E1 cells culture The coated AZ31 discs and the bare AZ31 control samples were placed into the 48-well plates (Corning, NY, USA). 1 mL of ␣-MEM supplemented with 10% FBS, 100 I.U./ml Penicillin and 100 ␮g/mL Streptomycin was added per well. Mouse osteoblastic MC3T3-E1 cells were cultured in a humidified atmosphere of 95% air and 5% CO2 at 37 ◦ C for 24 h and 72 h, respectively, at a seeding density of 4 × 104 cells/disc. Cell morphology was examined by SEM. After incubation for 1 day and 3 days, the discs were washed with PBS twice, fixed in 2.5% glutaraldehyde for 1 h, dehydrated with graded ethanol (30%,

50%, 75%, 95%, and 100% for 10 min, respectively), critical point dried, gold sputter coated and observed by SEM. Cell viability was determined by MTT assay kit (Sigma-Aldrich). After 1 and 3 days’ culture in 48-well plates, 0.5 mg/mL of MTT solution were added to each well. The plates were incubated for an additional 4 h at 37 ◦ C, and then the medium in each well was replaced by 500 ␮L of DMSO. Absorbance at a wavelength of 570 nm of the solution in each well was measured using a microplate reader (Biotek-Cytation 3). Each experiment was repeated independently at least three times. 3. Results and discussion 3.1. Characterization of CAp@PDA and Ag@PDA nanoparticles The phase and chemical composition of the hydrothermally synthesized biomimetic CAp nanoparticles (with carbonate content of 3.7 wt.%) have been identified in our previously work by XRD, Fourier transformed infrared spectroscopy (FTIR), ICP and thermogravimetric analysis (TGA) [32]. Here, we only show the TEM images of the CAp (Fig. S2a). The CAp nanocrystals are mainly in rectangular shape with average length of ∼75 nm and an apsect ratio of 2.5. There are a small amount of nanoparticles in hexagonal shape with side length of ∼30 nm. To verify that the PDA was coated onto the CAp, we also performed TEM (Fig. S2b) and high resolution TEM (Fig. S2c) analyses on CAp@PDA. High resolution TEM clearly reveals the interface between CAp nanocrystals and the PDA coating. A very thin layer of PDA (∼2 nm) was observed. Similarly, Fig. S3a shows the morphology of obtained Ag nanoparticles. The characteristic absorption wavelength at ∼ 440 nm in the UV–vis spectrum (Fig. S4) confirms that the obtained nanoparticles are silver. The size of the prepared Ag nanoparticles follows log-normal distribution with the average of ∼ 80 nm determined by DLS technique (Fig. S5). The Ag@PDA nanoparticles show typical core-shell features with the PDA layer of ∼ 25 nm thick (Fig. S3b). These results demonstrate that as expected, after immersion in a Tris-buffered solution containing dopamine under ultrasonication, both CAp and Ag nanocrystals were coated with a PDA layer despite with different thickness. The PDA seems easier to coat onto the Ag nanoparticles than the CAp nanoparticles, as reflected by their coating thickness. Such interesting phenomena is probably due to that PDA anchoring on the surface of materials is mainly by covalently bond and noncovalent interactions (e.g., ␲–␲ stacking, charge transfer, and hydrogen bonding) and the catechol group of dopamine performs stronger coordination and chelating ability to metal nanoparticle [27,34]. 3.2. Surface morphology, chemical composition and coating thickness We have exmined the surface morphology and chemical composition evolution after each step of surface treatment, as shown in Figs. 1 and 2, respectively. In contrast to the mechanically polished bare AZ31 with smooth surface (Fig. 1a), alkaline treatment makes the surface (Fig. 1b) much rough. The PDA coated AZ31 has (Fig. 1c) crazing surface probably caused by the corrosion of the Mg substrate during the self-polymerization of dopamine in the Tris buffer with pH = 8.5. The release of hydrogen makes it relatively hard to form uniform and compact PDA film. After grafting with chitosan, a dense surface was observed (Fig. 1d). As expected, CAp (Fig. 1e) and silver nanoparticles (Fig. 1f) were successfully immobilized onto the chitosan intermediate layer. The XPS spectra of the samples after each step of surface treatment were shown in Fig. 2 and Fig. S6. Compared to the wide scan survey spectrum of bare AZ31 sample (Fig. S6a), the presence

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Fig. 1. Surface morphology of the samples after each step of surface treatment. (a) mechanically polished AZ31, (b) alkaline pre-treated AZ31, (c) PDA coated AZ31, (d) chitosan grafted PDA, (e) chitosan immobilized with CAp@PDA nanoparticles, and (f) the final chitosan-CAp@PDA-chitosan-Ag@PDA multilayered coating.

of Na 1s at 1071.6 eV in the alkaline pre-treated sample (Fig. 2a) was observed. High resolution C 1s spectrum (Fig. 2b) reveals that the products after alkaline treatment are mainly composed of two components: (1) at 288.5 eV, which are assigned to sodium and magnsium carbonates; and (2) at 285.0 eV, which is normally interpreted as carbon in the form of C-C/C H groups. Such C C/C H groups usually appear on the outer surface (<3 nm in thickness) of almost all the metals in contact with the atmosphere at room temperature, irrespective of its composition [35,36]. The N 1 s peak at 399.3 eV (Fig. 2c) is charateristic for amine NH in the PDA structure [37]. Moreover, the C 1s core-level spectrum (Fig. 2d) can be fitted with four peaks at 284.6 eV, 285.6 eV, 286.2 eV and 287.9 eV, attributable to C H, C N, C O and C O groups of the PDA coating. These results confirm that a PDA transition layer was successfully deposited on the alkaline pre-treated AZ31 Mg alloy substrate. After grafted with chitosan, the intensity of C 1s core-level spectrum is significantly increased. High resolution C 1s spectrum further reveals that the intensity of C O peak dramatically increased which is due to the hydroxyl group in the backbone of chitosan. Also, the increase in the C O/N C O peak intensity was attributed to amide bonds formed between PDA and chitosan and the acetyl groups from chitosan backbone [38,39]. The XPS wide scan survery spectrum of the eventually obtained nano-multilayered coating (Fig. S6b) shows pronounced characteristic Ag peaks since the top layer is composed of Ag@PDA nanoparticles. The cross-section of the multilayered coating was shown in Fig. S7. The coating (between Pt and AZ31 alloy substrate) has a thickness of approximately 100 nm (Fig. 3).

obviously increased the corrosion potential to a more positive value from −1.57 V to −1.50 V and reduced the corrosion current density from 1.2 × 10−4 A/cm2 to 1.0 × 10−4 A/cm2 . The corroded surface morphology after polarization tests is shown in Fig. S8. The surface of bare AZ31 sample shows severe corrosion damage and extensive large cracks. In contrast, the coated sample shows only slight corrosion. All these results suggest that the fabricated nano-multilayered coating could improve the corrosion resistance of magnesium alloys. 3.4. In vitro degradation and Ag+ ions release The effect of the coating on the in vitro degradation behavior and the Ag+ ions release from the coating were investigated by immersing the samples in SBF for pre-defined intevals. The Mg2+ ions release profile (Fig. 5a) and pH evolution with time (Fig. 5b) demonstrate that the coating effectively reduce the amount of Mg2+ ions and lower the pH value of the SBF, which suggest that the nanomultilayered coating can slow down the degradation of magnesium alloys, in consistent with the above results from electrochemical potentiodynamic polarization tests. The cumulative release profile of Ag+ ions over 14 days from the coating is shown in Fig. 6. The coated AZ31 sample can release Ag+ ions in a sustained manner. During the first day of immersion, a burst release of Ag+ ions was observed with an amount of ∼ 300 ␮g/cm2 . Then, the release rate was dramatically decreased, at an average rate of ∼25 ␮g/cm2 . 3.5. Antibacterial activity

3.3. Electrochemical corrosion behavior The effect of coating on the corrosion resistance was investigated by electrochemical potentiodynamic polarization tests. Open circuit potential (OCP) evolving with time (1 h) for both bare and coated AZ31 Mg alloy discs in SBF are presented in Fig. 4a. The coated sample has much positive OCP value. Fig. 4b shows the typical potentiodynamic polarization curves of the two samples. The cathodic polarization curves represent the cathodic hydrogen evolution through water reduction, whereas the anodic curves represent the dissolution of Mg alloy. The extracted corrosion potential and corrosion current density from Tafel plots of both bare AZ31 and coated AZ31 samples show that the coating

Since the initial adhesion of bacteria onto a device surface is a critical step for pathogen colonization, biofilm formation, and subsequent device infection, it is of great importance to inhibit such bacteria adhesion at the early stage. A period of 4 h has been reported to be sufficient for bacteria to complete initial adhesion onto a surface [40,41]. Fig. 7 shows the representative optical photographs of E. Coli and S. aures bacterial colonies after incubation with bare and coated AZ31 for 2 h and 4 h, respectively. The number of bacterial colonies was quantified in Fig. 7. Compared to the bare AZ31, the number of bacterial colonies was significantly decreased for the coated AZ31. The antibacterial rates of the coated AZ31 with Ag@PDA as the outmost-layer against E. coli and S. aureus

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Fig. 2. XPS spectra of the samples of alkaline pre-treated AZ31 (a, b), PDA coated AZ31 (c, d) and chitosan grafed PDA (e, f). The left column shows the wide scan survey spectra and the right column shows the corresponding high resolution C 1s spectra.

reach 77.6% and 67.0% after 2 h, respectively, and reach 94.0% and 83.3% after 4 h, respectively. The antibacterial rates of the coated AZ31 with CAp@PDA as the outmost-layer against E. coli and S. aureus reach 80.0% and 57.8% after 2 h, respectively, and reach 83.3% and 91.0% after 4 h, respectively, indicating excellent antibacterial activity of the nano-multilayered coating. Such excellent antibacterial activity should be attributed to the synergetic effect of the Ag nanoparticles and chitosan, both of which are served as antimicrobial agents [33].

SEM micrographs of bacterial morphology (Fig. S9) show that the bacteria distrubted uniformly on the surface of bare AZ31 after incubation for 2 h. Much smaller number of bacteria was observed on the coated AZ31. Such observations are consistent with the results shown in Fig. 7. High magnification SEM images (Fig. S10) reveal the morphological changes of the individual bacteria. The morphogical observations further directly verifed that the multilayered coating has excellent antibacterial activity.

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Fig. 3. Open circuit potential (OCP) evolving with time and the typical potentiodynamic polarization curves of bare and coated AZ31 magnesium alloy discs with Ag@PDA as the outmost layer in simulated body fluid (SBF).

Fig. 4. The Mg2+ ions release profile (a) and pH evolution with time (b) in SBF.

Fig. 5. The cumulative release profile of Ag+ ions from the coating after immersion in SBF at 37 ◦ C for 14 days (n = 3).

3.6. In vitro cytocompatibility The osteoblast-like MC3T3-E1 cells proliferation after cultivation for 1 and 3 days on the bare and coated AZ31 Mg alloy were shown in Fig. 8. During the first day, the MC3T3-E1 prolifera-

tion/viability cultivated on the coated AZ31 with Ag@PDA as the outmost-layer was slightly lower than that of bare AZ31, which was probably due to the burst release of antibaterial agents at the early stage which may also have a negative effect on the cell proliferation, but the coated AZ31 with CAp@PDA as the outmost-layer has significantly higher cell proliferation/viability than that of bare AZ31, which may be attributed to the presence of CAp nanoparticles. However, after 3 days’ cultivation, both coated AZ31 samples demonstrate similar cell profileration/viability, which are much higher than that of bare AZ31. The effect of the nano-multilayered coating on the cell adhesion and spreading was examined by SEM after 1 and 3 days’ incubation. The typical morphology of MC3T3-E1 cells on the bare and coated AZ31 was shown in Fig. S11. The MC3T3-E1 cells show poor cell adhesion and spreading after incubation for 1 day on the surface of bare AZ31 and coated AZ31 with Ag@PDA as the outmostlayer. In contrast, the surface of the nano-multilayeed coating with CAp@PDA as the outmost-layer becomes more suitable for cell adhesion and spreading after 1 day’s incubation. However, the MC3T3-E1 cells show much less cell adhesion after incubation for 3 days on the surface of bare AZ31, but the nano-multilayered coating show good cell adhesion and spreading, as demonstrated by the increased amount and area of focal contacts of MC3T3-E1 cells.

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Fig. 6. The representative optical photographs of E. coli and S. aures bacterial colonies after incubation with bare and coated AZ31 for 2 h and 4 h, respectively.

Fig. 7. The statistical results of bacterial colonies after incubation with bare AZ31 and coated AZ31 for 2 h and 4 h, respectively (n = 3). (a) and (b) are the total colony forming units of E. coli and S. aureus in 1 mL medium (CFU/mL), respectively.

Both MTT assays and cell morphology observations suggest that by co-immobilization of CAp and Ag nanoparticles, the coating also demonstrate excellent in vitro cytocompatibility. 4. Conclusion In this study, we have developed a novel functional and biocompatible coating through dopamine and chitosan mediated layer-by-layer assembly of osteoinductive carbonated apatite and antibacterial silver nanoparticles, on the surface of AZ31 magnesium alloys. Electrochemical potentiodynamic polarization tests show that the fabricated nano-multilayered coating can increase

the corrosion potential to a more positive value and reduce the corrosion current density. Static immersion tests show that the coating can reduce the amount of Mg2+ ions, lower the pH value of the SBF and realize sustained release of the antibacterial Ag+ ions. Anteribacterial assays show that the numbers of bacterial colonies of both E. coli and S. aureus were significantly decreased for the coated AZ31. Osteoblastic MC3T3-E1 cells culture shows that the obtained coating also shows good in vitro cell cytocompatibility. In conclusion, the present nano-multilayered coating can not only enhance the corrosion resistance and therefore, slow down the biodegradation, but also significantly increase the antibacterial activity and enhance the biocompatility of magnesium alloys.

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Fig. 8. The results of MTT assay. The optical density was read at a wavelength of 570 nm after incubation for 1 and 3 d on each group of samples. Error bars represent ± standard deviation (n = 3, p < 0.05).

Acknowledgements This work was financially supported by the National Key R & D Program of China (2016YFB0700803), the Natural Science Foundation of Guangdong Province (Grant No. 2014A030310358) and the Special Funds for Fundamental Research of Strategic Emerging Industries of Shenzhen (Grant No. JCYJ20150331101823688). Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.colsurfb.2017.06. 013. References [1] M.P. Staiger, A.M. Pietak, J. Huadmai, G. Dias, Magnesium and its alloys as orthopedic biomaterials: a review, Biomaterials 27 (2006) 1728–1734. [2] Y. Chen, Z. Xu, C. Smith, J. Sankar, Recent advances on the development of magnesium alloys for biodegradable implants, Acta Biomater. 10 (2014) 4561–4573. [3] F. Witte, N. Hort, C. Vogt, S. Cohen, K.U. Kainer, R. Willumeit, F. Feyerabend, Degradable biomaterials based on magnesium corrosion, Curr. Opin. Solid State Mater. Sci. 12 (2008) 63–72. [4] S. Zhang, X. Zhang, C. Zhao, J. Li, Y. Song, C. Xie, H. Tao, Y. Zhang, Y. He, Y. Jiang, Y. Bian, Research on an Mg–Zn alloy as a degradable biomaterial, Acta Biomater. 6 (2010) 626–640. [5] L. Xu, F. Pan, G. Yu, L. Yang, E. Zhang, K. Yang, In vitro and in vivo evaluation of the surface bioactivity of a calcium phosphate coated magnesium alloy, Biomaterials 30 (2009) 1512–1523. [6] H.M. Wong, K.W.K. Yeung, K.O. Lam, V. Tam, P.K. Chu, K.D.K. Luk, K.M.C. Cheung, A biodegradable polymer-based coating to control the performance of magnesium alloy orthopaedic implants, Biomaterials 31 (2010) 2084–2096. [7] J. Li, Y. Song, S. Zhang, C. Zhao, F. Zhang, X. Zhang, L. Cao, Q. Fan, T. Tang, In vitro responses of human bone marrow stromal cells to a fluoridated hydroxyapatite coated biodegradable Mg–Zn alloy, Biomaterials 31 (2010) 5782–5788. [8] S.V. Dorozhkin, Calcium orthophosphate coatings on magnesium and its biodegradable alloys, Acta Biomater. 10 (2014) 2919–2934. [9] G. Wu, J.M. Ibrahim, P.K. Chu, Surface design of biodegradable magnesium alloys—a review, Surf. Coat. Technol. 233 (2013) 2–12. [10] H.M. Wong, Y. Zhao, V. Tam, S. Wu, P.K. Chu, Y. Zheng, M.K.T. To, F.K.L. Leung, K.D.K. Luk, K.M.C. Cheung, K.W.K. Yeung, In vivo stimulation of bone formation by aluminum and oxygen plasma surface-modified magnesium implants, Biomaterials 34 (2013) 9863–9876. [11] C.D. Gu, W. Yan, J.L. Zhang, J.P. Tu, Corrosion resistance of AZ31 B magnesium alloy with a conversion coating produced from a choline chloride—Urea based deep eutectic solvent, Corros. Sci. 106 (2016) 108–116. [12] J. Zhang, C. Gu, J. Tu, Robust slippery coating with superior corrosion resistance and anti-icing performance for AZ31B Mg alloy protection, ACS Appl. Mater. Interfaces 9 (2017) 11247–11257. [13] X.N. Gu, W. Zheng, Y. Cheng, Y.F. Zheng, A study on alkaline heat treated Mg-Ca alloy for the control of the biocorrosion rate, Acta Biomater. 5 (2009) 2790–2799.

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