Corrosion resistance and antibacterial properties of hydroxyapatite coating induced by gentamicin-loaded polymeric multilayers on magnesium alloys

Accepted Manuscript Title: Corrosion resistance and antibacterial properties of hydroxyapatite coating induced by gentamicin-loaded polymeric multilayers on magnesium alloys Authors: Xiao-Jing Ji, Ling Gao, Jia-Cheng Liu, Jing Wang, Qiang Cheng, Jian-Peng Li, Shuo-Qi Li, Ke-Qian Zhi, Rong-Chang Zeng, Zhen-Lin Wang PII: DOI: Reference:

S0927-7765(19)30252-8 https://doi.org/10.1016/j.colsurfb.2019.04.029 COLSUB 10153

To appear in:

Colloids and Surfaces B: Biointerfaces

Received date: Revised date: Accepted date:

30 January 2019 12 April 2019 13 April 2019

Please cite this article as: Ji X-Jing, Gao L, Liu J-Cheng, Wang J, Cheng Q, Li J-Peng, Li S-Qi, Zhi K-Qian, Zeng R-Chang, Wang Z-Lin, Corrosion resistance and antibacterial properties of hydroxyapatite coating induced by gentamicin-loaded polymeric multilayers on magnesium alloys, Colloids and Surfaces B: Biointerfaces (2019), https://doi.org/10.1016/j.colsurfb.2019.04.029 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

Corrosion resistance and antibacterial properties of hydroxyapatite coating induced by gentamicin-loaded polymeric multilayers on magnesium alloys Xiao-Jing Jia,1, Ling Gaob,c,1, Jia-Cheng Liub,c,d, Jing Wanga, Qiang Chenga, Jian-Peng Lia, Shuo-Qi Lia,*, Ke-Qian Zhib,c,*, Rong-Chang Zenga,*, Zhen-Lin Wange College of Materials Science and Engineering, Shandong University of Science and

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a

Technology, Qingdao, 266590, China

Department of Oral and Maxillofacial Surgery, the Affiliated Hospital of Qingdao University,

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Qingdao, Shandong, 266000, China c

Key Lab of Oral Clinical Medicine, the Affiliated Hospital of Qingdao University, Qingdao,

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Shandong, 266000, China

School of Stomatology, Qingdao University, Qingdao, Shandong, 266071, China

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College of Materials Science and Engineering, Chongqing University of Technology,

Chongqing, 400065, China Corresponding authors

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*

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d

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E-mail addresses: [email protected] (S.-Q. Li), [email protected] (K.-Q. Zhi) [email protected] (R.-C. Zeng). Both authors contribute equally to this work.

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Graphical abstract

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Highlights

The multilayers loaded with GS could be formed on Mg alloys by LbL assembly.

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The multilayers could promote the hydrothermal fabrication of HAp coating.

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The obtained (PAA/GS)20/PAA-HAp could decrease the corrosion rate of Mg alloys.

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The coating showed antibacterial property and a prolonged release profile for GS.

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This functional coating showed acceptable biocompatibility to MC3T3-E1 osteoblasts.

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Abstract As a result of their good biocompatibility, bioactivity, and mechanical properties, magnesium (Mg) alloys have received considerable attention as next generation biodegradable implants. 2

Herein, in order to achieve a proper degradation rate and good antibacterial ability, we reported a novel hydroxyapatite coating induced by gentamicin (GS)-loaded polymeric multilayers for the surface treatment of the Mg alloy. The coating was characterized by X-ray diffraction, fourier transform infrared spectroscopy and scanning electron microscopy. The

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as-prepared hydroxyapatite coating showed the compact morphology and a well-crystallized apatite structure. This coating could improve the adhesion strength and reduce the corrosion

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rate of the substrate in simulated body fluid solution. Meanwhile, the drug release and antibacterial experiments demonstrated that the GS loaded specimen revealed a significant antimicrobial performance toward Staphylococcus aureus and had a prolonged release profile

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of GS, which would be helpful to the long-term bactericidal activity of the Mg implant. This

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coating showed acceptable biocompatibility via MTT assay and Live/dead staining. Thus, the

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multilayers-hydroxyapatite coated Mg alloy could improve the corrosion resistance and biocompatibility while delivering vital drugs to the site of implantation. Magnesium

alloy,

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Keywords:

Corrosion

resistance,

Antibacterial

performance,

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Hydroxyapatite coating, Layer-by-layer assembly

1. Introduction

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Considering their similar elastic modulus and mechanical properties to natural bone,

magnesium (Mg) alloys are promising candidates as orthopedic implants in biomedical applications [1-4]. Nevertheless, the wide spread applications of Mg alloys are restricted due to the poor corrosion resistance [5-9]. To modify Mg implants for improved biological activity and anti-corrosion properties, hydroxyapatite (Ca10(PO4)6(OH)2, HAp) with excellent 3

biocompatibility and bioactivity, has been extensively established as useful coatings [10]. And in recent years, various synthetic strategies have been applied for the preparation of HAp coating, such as template method [11], spray drying [12] and hydrothermal synthesis [13]. Among the diverse methods available for deposition of HAp coatings, hydrothermal treatment

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has received considerable attention due to its simplicity and energy efficiency, and the obtained HAp coated Mg alloys could be used as implants in biomedical application,

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especially for repair and reconstruction of damaged bone tissues [14].

However, proteins, amino acids and other organic substances are easily adsorbed on the HAp surface, favoring the adhesion and growth of the bacterial and thus resulting

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implant-associated infections [15]. Although all the implant devices and operation medical

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apparatus should be strictly disinfected before implantation surgery, the possibility of the

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colonization of the bacterial still exists. Once a biofilm is formed on the implant surface, the bacteria will increase the corrosion rate and have strong drug resistance. Hence, together with

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corrosion protection, prevention of inflammatory reaction is also of great importance for

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surgical procedures based on HAp coated Mg alloys implants. Incorporating antibiotics in HAp coatings is one of the most effective approaches to

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alleviate this problem [16]. The widely used gentamicin sulfate (GS) is a broad-spectrum aminoglycoside antibiotic. It could be incorporated in bioactive apatite coatings according to

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various methods, and its sustained release from the GS-apatite composite could inhibit significant bacterial adhesion and prevented biofilm formation [17, 18]. Montemor et al. prepared the polycaprolactone/levofloxacin/nanohydroxyapatite composite coating with the desired biocompatible functionality and corrosion protection of the bare alloy [19]. However, this coating showed a porous and smooth morphology, resulting in the initial 45 % 4

burst release of antibiotics during the first 4 h. In fact, the release of antibiotics should be controlled in a consistent manner to avoid subsequent probable infections [20]. Therefore, it is better to develop a method whereby the antibacterial implants containing GS has anti-corrosion properties and a prolonged release profile.

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Nowadays, layer-by-layer (LbL) assembly is a versatile approach that has been employed in local drug delivery system and bioactive coatings for antibacterial application

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[21]. At the same time, the LbL assembled multilayer films could also affect the surface morphology and the thickness of HAp coatings, which can mitigate the burst release of GS from the dense and thick HAp coating [22]. Polyelectrolytes were often used as building

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blocks in the LbL process and were chosen to be the counterpart of GS in this research to

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fabricate the multilayers. As a polyelectrolyte with abundant carboxylate groups, polyacrylic

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acid (PAA) is highly anionic and acidic, which enables it to be an appropriate unit with negative charge in LbL system driven by electrostatic forces. Meanwhile, PAA is also a

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water-soluble polymer with a high binding capacity, which would contribute to the formation

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of Ca-P coating due to the interaction between −COO− and Ca2+ ions [23]. Thus, a dense and compact HAp coating was constructed on the multilayers coated specimen by hydrothermal

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treatment. In this work, multilayer-induced HAp coating was conducted sequentially to modify AZ31 alloys and their corrosion resistance, antimicrobial and cytocompatibility

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characteristics were determined systematically. 2. Experimental 2.1. Materials and chemicals The experimental material was an as-extruded Mg alloy AZ31 (Al 2.5-3.0, Zn 0.7-1.3, Mn > 0.2 and the balance Mg), which was supplied by the Shandong Yinguang Yuyuan Light 5

Metal Precise Forming Co., Ltd., China. PAA (MW = 800-1000) was purchased from Qingdao Jingke Chemical Reagent Co., Ltd., China. Poly(ethyleneimine) (PEI, MW = 600) and GS (MW = 575.67) were purchased from Shandong Xiya Chemical Technology Co., Ltd., China. All the regents were of analytical grade and ultra-pure water was used throughout the

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experiment. 2.2. Alkaline treatment of the substrate

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The AZ31 alloys were cut into rectangular shapes with 20 × 20 × 5 mm dimensions for surface characterization. While for the biological experiment, the specimens were cut to the size of 10 × 10 × 3 mm. All the substrates were ground with SiC papers ranging from a grit

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size of 150-2500, washed with water and ethanol, and then dried with warm air. The polished

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substrates were soaked in 5 mol L-1 NaOH solution for 30 min, followed by thoroughly

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cleaning with ultra-pure water and drying with warm air. 2.3. Preparation of composite coatings

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First, the NaOH-treated substrates were immersd into the PEI solution (5 mg mL-1, pH =

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10.0) for 30 min. Then, PAA (1 mg mL-1, pH = 7.0) and GS (1 mg mL-1) were deposited alternatively on the sample by dip-LbL assembly and washed with water. The multilayer

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coatings could be obtained by repeating this process for 20 times [24]. To make the outmost layer negatively charged, the samples were treated with the PAA solution at the end of the

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assembly process and finally, the samples were dried at 80 °C for 30 min to obtain the (PAA/GS)20/PAA multilayers coated Mg alloys. For the fabrication of the HAp coating, the samples were perpendicularly mounted into the teflon-lined stainless steel autoclaves of 100 mL capacity which contained the hydrothermal solution consisted of 14 mmol L-1 Ca(NO3)2, 8.4 mmol L-1 NaH2PO4 and 4 6

mmol L-1 NaHCO3. Afterwards, the stainless reactors were sealed and kept at 150 °C for 4 h. After the vessel was cooled to room temperature, the samples were rinsed with water and dried with warm air. The obtained coating was referred as (PAA/GS)20/PAA-HAp. Fig. 1 showed the schematic representation of the surface modification of the Mg alloy. For

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comparison, the pure HAp coated samples were also prepared without the LbL process. 2.4. Surface analysis

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The surface microstructure of the samples was characterized by scanning electron microscopy (SEM, Nova Nano SEM 450, FEI Corporation, USA) equipped with energy-dispersive X-ray spectroscopy (EDS). The phase compositions were identified by

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X-ray diffraction (XRD, Rigaku D/MAX 2500 PC, Rigaku Corporation, Japan). Fourier

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transform infrared spectroscopy (FT-IR, Nicolet 380, Thermo Electron Corporation, USA)

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was used to examine the surface chemical structures of the samples. The adhesion strength of the coatings was evaluated by nano-testing machine (Nano Test 600, Micro Materials, UK)

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with a Rockwell diamond probe.

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2.5. Corrosion characterization

The corrosion measurements were performed using electrochemical potentiodynamic

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polarization (PDP), electrochemical impedance spectra (EIS) and hydrogen evolution (HE) test in stimulated body fluid (SBF). The detailed preparation procedure of SBF and the

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electrochemical test system were referred to our previous publication [24]. 2.6. Antibacterial assays The antibacterial activity of the samples was tested against gram-positive Staphylococcus aureus (S. aureus, ATCC 6538) by plate-counting method. The samples were placed in the tubes, covered with 3 mL of broth containing approximately 106 colony forming units per 7

milliliter (CFU mL-1) concentration of S. aureus and incubated by rotation for 4 h at 37 °C [25]. After incubation, 0.1 mL of the suspension was diluted with 0.8 % NaCl 103-fold, and then 0.1 ml of the diluted bacteria solution was spread evenly onto a nutrient agar plate. Subsequently, all the plates were incubated at 37 °C for 24 h and the number of colonies

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formed on the agar plate was counted. Each treatment was carried out in triplicate. 2.7. Loading and releasing of GS

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The procedure for the loading dosage calculation of GS was adapted from Ref. [26]. First, we have prepared the solutions with various GS concentrations to obtain the absorption spectra of GS by UV–vis spectrophotometry (UV-3101PC, Shimadzu Corporation, Japan).

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Then the maximum absorption wavelength λmax = 247 nm was selected to draw the standard

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curve of GS, which could be used to get the exact concentration of the drug solution. By

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deducting the GS amount in the remaining solution and washing solution from the total GS amount, the GS loading dosage could be obtained. The experiment was performed in triplicate

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for (PAA/GS)20/PAA-HAp coating.

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To quantify the release of GS, (PAA/GS)20/PAA-HAp coated samples were immersed in 8 mL phosphate buffer solutions (PBS, pH = 7.4) under agitation (140 rpm) at 37 °C for 384 h.

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Then 4 mL drug release medium was withdrawn at predetermined time intervals, and replaced with 4 mL fresh PBS at each measurement to continue the in vitro release study. The

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concentration of GS was measured by UV–vis spectrophotometry. For the drug-loaded sample, the experiment was conducted in triplicate. 2.8. Cytocompatibility tests 2.8.1. Cell proliferation assay Mouse MC3T3-E1 pre-osteoblasts were selected for the in vitro cell culture tests. Cells 8

were cultivated in alpha Dulbecco's modified Eagle's medium (α-DMEM) supplemented with 10 % foetal bovine serum (FBS) and 1 % penicillin-streptomycin solution in a humidified atmosphere with 5 % CO2 at 37 ℃. Indirect contact method of culturing the cells with the different extracts was used to assess the biocompatibility. Prior to the experiments, all the

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samples were sterilized at high temperature and pressure. After that, the extracts were prepared with a sample surface area to extraction medium

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(α-DMEM) ratio of 1 mL cm-2 at 37 ℃ in a humidified atmosphere with 5 % CO2 for 3 days. The supernatant fluids were collected and stored at 4 ℃. The seeded cells were cultured on a 96-well plate at a density of 3000 cells per well for 24 h to allow attachment. Then, the

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medium was replaced by 100 µL of α-DMEM supplemented with 20% extracts and 10 % FBS.

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After culturing for 1, 3 and 5 days, the extracts were replaced by the fresh medium and 10 µL

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of 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT, M8180, Solarbio, China) solution was added to each well. After incubation for 4 h, the blue formazen crystals

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were formed and were dissolved by dimethyl sulfoxide (DMSO). The absorbance was

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measured at a wavelength of 492 nm by a microplate reader (Multiskan Go 1510, Thermo Fisher Scientific, China). The cell viability was calculated as follows: (%) = (OD

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Cell viability

sample

/OD

control

(1)

) × 100

where ODsample and ODcontrol represent the optical density (OD) values of cells cultured with

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the sample and without the sample, respectively. 2.8.2. LIVE/DEAD viability assay Osteoblasts were cultured on a 24-well plate at a density of 5000 cells per well. After 6 h, the cell culture media in each well were replaced by 500 µL mixed media with 20 vol% extracts and 80 vol% regular cell culture media. Then, cells were cultured in a humidified 9

atmosphere of 5 % CO2 at 37 ℃. After incubation for 24 h, the cells were stained by LIVE/DEAD® assay (L3224, Invitrogen, China). Briefly, the mixed solution of Calcein-AM and Ethidium homodimer was added to each well and incubated for 30 min. After that, the solution was removed and the cells were observed by a fluorescence microscope (EVOS,

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Invitrogen, China). The living and dead cells were stained by Calcein-AM and Ethidium homodimer into green and red colors, respectively. The cells cultured in fresh media without

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samples were used as negative controls and each sample had three replicas. 2.9. Statistical analysis

Data of the biocompatibility tests were reported as the mean ± standard deviation (SD).

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Statistical analysis was performed by the independent-samples t-test. Differences were

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considered significant at *P < 0.05 or **P < 0.01. All the experiments were carried out three

3. Results and discussion

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times, with three replicates used for each test.

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3.1. Composition and surface characterization of the coatings

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SEM was used to analyze the morphology of the different HAp coatings (Fig. 2). The pure HAp coating (Fig. 2a and c) presented a distinctive bilayer architecture: the up layer

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consisted of flower-like HAp crystals and the bottom layer comprised loose grass-like HAp crystals, while (PAA/GS)20/PAA-HAp coating (Fig. 2b and d) exhibited a compact flake

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structure. Additionally, from the EDS spectra (Fig. 2e and f), the Ca/P ratios were calculated to be 1.40 for pure HAp coating and 1.84 for (PAA/GS)20/PAA-HAp coating. The results indicated that pure HAp coating was calcium-deficient, resulting from the substitutions of Ca2+ ions by small amounts of Mg2+ ions in the HAp lattice [27]. However, the higher Ca/P ratio for (PAA/GS)20/PAA-HAp coating suggested that more Ca2+ ions were incorporated into 10

the HAp crystal lattices, which played a positive role in the formation of HAp [28]. The strong interaction between carboxyl groups of PAA and Ca2+ ions had been well documented, and thus Ca2+ ions could be absorbed and quickly concentrated on the surface with abundant carboxyl groups, resulting the promotion of the fabrication of the calcium-rich apatite [29].

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From the cross-sectional SEM and EDS line scan images (Fig. S1), we could see that the thickness of pure HAp coating (Fig. S1a and c) was 11.26 μm, while (PAA/GS)20/PAA-HAp

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coating (Fig. S1b and d) consisted of densely packed flake-shape precipitates and had the higher thickness of 21.82 μm. It was also showed that (PAA/GS)20/PAA-HAp coating was closely adherent to the substrate with no gaps or cracks between the coating and Mg substrate.

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We thus speculated that the formation of a relatively thick and flawless HAp layer might not

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only improve the corrosion resistance but also lead to the prolonging release of GS due to the

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elongated diffusion path and compact structure [30, 31]. Moreover, the denser surface topography of (PAA/GS)20/PAA-HAp coating might lead to

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higher adhesive strength. The results of scratch test showed that the critical loads were 1864

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mN for pure HAp coating and 2758 mN for (PAA/GS)20/PAA-HAp coating (Fig. S2), respectively, which indicated that the adhesion strength of the coating prepared by LbL

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assembly and further hydrothermal treatment was stronger than the coating via direct hydrothermal treatment, which could be due to the flawless and dense structure induced by

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the multilayers.

XRD patterns of pure HAp and (PAA/GS)20/PAA-HAp coatings were presented in Fig.

3a. The results indicated that, besides the diffraction peaks from AZ31 substrate, the characteristic peaks of Mg(OH)2 and HAp were detected, demonstrating that Mg(OH)2 and HAp were produced as the main components of the coatings on AZ31 alloy. For both the HAp 11

coatings, the characteristic peaks of HAp appeared at 2θ ≈ 10.7 °, 21.7 °, 22.8 °, 25.8 °, 28.2 °, 31.8 °, 49.4 °, 50.6 °, 53.8 °, which confirmed the presence of well-crystallized HAp phase [32]. The HAp coating with high crystallinity would have strong binding force and could decrease the degradation rate of the HAp coating, thus further prolonging the service life span

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of implants [33]. Therefore, it could be speculated that both of the two coatings might reduce the corrosion rate of Mg alloys.

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FT-IR was also used to confirm the successful fabrication of HAp coatings (Fig. 3b). For the two coatings, we could observe the typical bands such as fundamental vibrational modes at around 1097, 1036, 604 and 563 cm−1 for PO43− groups [23], Mg-OH vibration mode at

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3705 cm−1 for Mg(OH)2 [10], the stretching and vibration modes at 3570 cm−1 for O-H groups

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[24] and vibrational mode at 871 cm−1 for CO32– groups [34]. These peaks could thus clearly

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confirm the formation of typical hydroxyapatite and small content of carbonate-substituted hydroxyapatite [35]. Besides, the main differences between the two coatings were visible in

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N–H band at 1569 cm−1 and -COO- band at 1453 cm−1 [36, 37]. These bands were present

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only in (PAA/GS)20/PAA-HAp coating, indicating the successful loading of GS and the presence of PAA, in which the carboxylic groups were dissociated into COO− forms [23].

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3.2. Corrosion behavior

Theoretically, the lower corrosion density (Icorr) in PDP curves revealed the better

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corrosion resistance. In Fig. 4a, (PAA/GS)20/PAA-HAp coating (Icorr= 8.58 × 10-7 A cm-2) exhibited higher corrosion resistance than pure HAp coating (Icorr= 9.44 × 10-6 A cm-2) and AZ31 substrate (Icorr= 3.52 × 10-5 A cm-2). This increased corrosion protection performance could be due to the dense and thick surface coverage of the HAp, as proven by the SEM and XRD [38]. 12

EIS measurements were also conducted. In the Bode plots, it could be seen that (PAA/GS)20/PAA-HAp coating showed the highest impedance modulus |Z| value (9944.64 Ω cm2) at 0.01 Hz (Fig. 4b), which exhibited the best anti-corrosion property of this coating [39]. Moreover, the phase angle of (PAA/GS)20/PAA-HAp coating became loftier and wider (Fig.

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4c), revealing that the coating was uniform and dense [24]. At the same time, Nyquist plots and the corresponding electrical equivalent circuits (EECs) were obtained (Fig. 4d-f). The

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EECs contained Rs, Rc, Rct, RL, L, C and CPE, corresponding to the resistance of solution, the resistance of coatings, the charge transfer resistance, inductive resistance, inductance, the double layer capacitance, constant phase elements (CPEs) respectively. By calculation, values

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of Rct up to 4654 Ω cm2 had been recorded for (PAA/GS)20/PAA-HAp coating, as against

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1555 Ω cm2 and 66 Ω cm2 obtained in pure HAp coating and AZ31 substrate, respectively.

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Generally, a higher Rct value implied the better anti-corrosion property of the samples [8]. Thus, these results showed that (PAA/GS)20/PAA-HAp coating had the best corrosion

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resistance, which was also in accordance with the electrochemical test results.

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The HER results of different samples immersed in SBF solution for 160 h were displayed in Fig. 5a. The HER for the samples could be ranked in increasing order as follows:

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(PAA/GS)20/PAA-HAp coating (0.026 ± 0.007 mL cm−2 day−1) < HAp coating (0.259 ± 0.031 mL cm−2 day−1) < AZ31 substrate (0.492 ± 0.041 mL cm−2 day−1). Compared to the bare AZ31

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and pure HAp coating, (PAA/GS)20/PAA-HAp coating showed the lowest HER, which was in agreement with the electrochemical data. The pH changes during corrosion as a function of time for 160 h were shown in Fig. 5b. It could be found that the pH values of bare Mg alloy and pure HAp coating increased rapidly at the initial stage of immersion and then kept nearly stable at about 7.82 and 7.63, 13

respectively, which might be due to the alkalization of the medium [40]. However, for (PAA/GS)20/PAA-HAp, the pH of SBF solution after 160 h of immersion lied around 7.40, confirming that the coating could provide sufficient protection for Mg substrates from the aggressive attack by the corrosive media.

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3.3. Immersion tests From Fig. 6a and d, it could be seen that the bare Mg alloy suffered obvious corrosion

the

surface

layer

had

peeled

off

due

to

the

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and large cracks could be observed on the surface. While for pure HAp coating (Fig. 6b and e), hydrogen

evolution.

However,

(PAA/GS)20/PAA-HAp coating (Fig. 6c and f) remained a relatively intact surface

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characteristic and no obvious corrosion pits could be found. From the EDS data, we could

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also see the Ca/P atomic ratio was 1.68 for the (PAA/GS)20/PAA-HAp coating (Fig. S3),

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which was slightly lower than the value of the LbL-induced HAp coating before the immersion. Obviously, due to the increasing OH− concentration caused by the release of Mg2+,

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the pH values would increase for the (PAA/GS)20/PAA-HAp coated samples (Fig. 5b). And

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then, the released Mg2+ ions might prevent the crystallization of HAp with the substitution of Mg2+ ions for the Ca2+ ions in the HAp structure, resulting in the lower Ca/P ratio of

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(PAA/GS)20/PAA-HAp coating after an immersion. 3.4. In vitro antibacterial activity

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Fig. 7a and b showed the number of bacterial colonies on the agar plates of different

samples. The CFUs of S. aureus increased in the following order: (PAA/GS)20/PAA-HAp coating (1 ± 1 CFU) < AZ31 substrate (538 ± 16 CFU) < pure HAp coating (1659 ± 24 CFU) < blank group (1913 ± 30 CFU). On the whole, (PAA/GS)20/PAA-HAp coated Mg alloys possessed effective antibacterial properties due to the local high GS concentration, promising 14

to be used as antimicrobic material to prevent the undesirable accumulation of bacteria on substrate surfaces. 3.5. Loading and release of GS from (PAA/GS)20/PAA-HAp coating As shown in Fig. 7c, the GS loading dosage was 83.33 ± 15.55 µg cm-2 for 5 bilayers of

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(PAA/GS)n/PAA multilayer films and it steadily increased into 331.56 ± 22.83 µg cm-2 as the assembled number increased to 20. This result showed that the GS loading amount for each

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bilayer was about 16.58 µg cm-2 and the total loading dosage could be easily tuned by adjusting the number of assembled layers. It was reported that the loading dosage was 166.93 ± 16.38 μg cm-2 for the GS-loaded (PAA/(polyvinyl pyrrolidone/chitosan))10 multilayers by

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Chen et al. [41], and Hammond et al. also prepared the multilayer films containing

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poly(β-amino ester), hyaluronic acid and GS, which had the GS loading capacity of 122.97 μg

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cm-2 [42]. Thus, the relatively high loading dosage of GS via LbL with PAA was successfully achieved in this work.

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The release curves of GS from (PAA/GS)20/PAA-HAp coating were measured via

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UV-vis spectrophotometer, as showed in Fig. 7d. Obviously, (PAA/GS)20/PAA-HAp coating displayed desired drug release profile: burst release during the initial 24 hours (approximately

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5.2 µg cm-2, 23.1 %) and subsequent long-term slow release for 15 days (almost 19.98 µg cm-2, 88.78 %). Guo et al. prepared the GS-loaded carbonated hydroxyapatite coatings which

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could prolong the antibiotic release up to 24 hours [43]. Saeed Hesaraki et al. developed the poor crystalline carbonated hydroxyapatite submicron particles loaded with GS which could increase the release time till nearly 5 days [44]. Obviously, (PAA/GS)20/PAA-HAp coating developed in this work exhibited a much smaller sustained release rate. We speculated that this sustained release profile could be ascribed to the fact that the electrostatic attraction 15

between PAA and GS could decrease the diffusion and release rate of the drug [45]. Moreover, the dense and thick HAp coating could also act as physical barriers against the release of GS [31]. Thus, (PAA/GS)20/PAA-HAp coated Mg alloy could emerge as a promising material with drug delivery system for the bone implant replacement.

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3.6. Cytocompatibility tests The response of MC3T3-E1 cells was employed to check the cell viability in the

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presence of 20 % Mg alloys extracts. Fig. 8a and b showed the OD values and cell viability after incubating pre-osteoblasts for 1, 3 and 5 days. Obviously, when cultured for 1 day, the (PAA/GS)20/PAA-HAp group slightly reduced cell viability, which could be due to the burst

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release of GS in the first 24 h [25]. While after culturing for 3 and 5 days,

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(PAA/GS)20/PAA-HAp coated Mg showed enhanced osteogenesis effect, which could be

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ascribed to the much slower release rate of GS. At the same time, Live/Dead staining of osteoblasts 24 h post seeding could be seen in Fig. 8c-e. It was worth mentioning that the cells

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for all samples had generally healthy fusiform-like shape, and were wide spread with respect

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to morphology. Thus, it could be concluded that (PAA/GS)20/PAA-HAp group could display acceptable and enhanced cytocompatibility to osteoblasts.

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4. Conclusions

A novel (PAA/GS)20/PAA-HAp coating was successfully fabricated on Mg alloys

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through the dip-LbL technique and hydrothermal treatment. The drug-loaded multilayers were constructed by LbL deposition of negatively charged PAA and positively charged GS on the substrate via electrostatic interactions. The obtained (PAA/GS)20/PAA-HAp coating showed not only the compact flake morphology, but also an apatite structure with high crystallinity, which might be due to the interaction between carboxyl groups of PAA and Ca2+ ions. 16

Compared to the pure HAp coating, (PAA/GS)20/PAA-HAp coating exhibited denser and thicker structure, which could provide excellent corrosion protection for Mg alloys. Besides, this GS loaded coating presented antibacterial properties toward S. aureus and a prolonged release profile for the antibiotics (16 days). Moreover, (PAA/GS)20/PAA-HAp coated Mg

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alloys displayed acceptable biocompatibility to MC3T3-E1 osteoblasts. Thus, owing to the comprehensive exhibition of favourable corrosion resistance, controlled GS release, excellent

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antibacterial activity and suitable cytocompatibility, the (PAA/GS)20/PAA-HAp coating is attractive for applications in the surface modification of Mg implant in orthopaedic applications.

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Acknowledgments

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This work was supported by the National Natural Science Foundation of China (No.

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51571134), the Shandong University of Science and Technology (SDUST) Research Fund

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ZR2017BEM002).

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(No. 2014TDJH104) and the Shandong Provincial Natural Science Foundation (No.

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Figure captions

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Fig. 1. Schematic representation of the preparation of (PAA/GS)20/PAA-HAp coating.

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(PAA/GS)20/PAA-HAp (b, d, f) coatings.

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Fig. 2. SEM images (a-d) and EDS spectra (e, f) of pure HAp (a, c, e) and

Fig. 3. XRD patterns (a) of pure HAp and (PAA/GS)20/PAA-HAp coatings; FT-IR spectra (b) of PAA, GS, pure HAp and (PAA/GS)20/PAA-HAp coatings. 24

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Fig. 4. PDP curves (a), Bode magnitude plots (b) and Bode phase angle plots (c) of (i) AZ31

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substrate, (ii) pure HAp and (iii) (PAA/GS)20/PAA-HAp coatings; Nyquist plots and fitting

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curves and the corresponding electrical equivalent circuit models (d-f) of all samples in SBF

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solution.

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Fig. 5. HER results (a) and pH values (b) of (i) AZ31 substrate, (ii) pure HAp and (iii)

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(PAA/GS)20/PAA-HAp coatings in SBF solution for 160 h.

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Fig. 6. Digital camera photographs (a-c) and SEM morphologies (d-f) of (a, d) AZ31 substrate,

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(b, e) pure HAp and (c, f) (PAA/GS)20/PAA-HAp coatings after an immersion of 160 h.

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Fig. 7. CFU numbers of S. aureus colonies (a) and representative images (b) on different samples of (i) blank, (ii) AZ31 substrate, (iii) pure HAp and (iv) (PAA/GS)20/PAA-HAp coatings after culturing for 24 h; changes of amount of GS loaded in (PAA/GS)20/PAA multilayers with the increase of assembling cycles (c); the cumulative release concentration

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from (PAA/GS)20/PAA-HAp coating in PBS for 384 h (d).

Fig. 8. OD values (a) and Cell viability (b) of MC3T3-E1 cultured in different extracts prepared with negative control, AZ31 substrate and (PAA/GS)20/PAA-HAp for 1, 3 and 5 days. Statistically significant differences (*p < 0.05, **p < 0.01.); Fluorescent images (c-e) of 27

MC3T3-E1 after culturing for 24 h in extracts of the (c) negative control, (d) AZ31 substrate

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and (e) (PAA/GS)20/PAA-HAp.

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