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Cite this: DOI: 10.1039/c8bm00298c

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Layered double hydroxide/poly-dopamine composite coating with surface heparinization on Mg alloys: improved anticorrosion, endothelialization and hemocompatibility† Hua Li,‡a Feng Peng,‡b,c Donghui Wang,b Yuqin Qiao,b Demin Xu*a and Xuanyong Liu *b Magnesium (Mg) and its alloys are promising cardiovascular stent materials due to their favourable physical properties and complete degradation in vivo. However, rapid degradation and poor cytocompatibility hinder their clinical applications. To enhance the corrosion resistance and endothelialization of the AZ31 alloy, a layered double hydroxide (LDH)/poly-dopamine (PDA) composite coating (LDH/PDA) was successfully fabricated. Polarization curves and the electrochemical impedance spectroscopy Nyquist spectrum test proved that the corrosion resistance of the LDH/PDA sample was significantly improved in vitro. The LDH/PDA sample greatly improved the adherence process and the proliferation rate of human umbilical vein endothelial cells (HUVECs). After culturing for 10 days, the number of living HUVECs on the LDH/ PDA sample was comparable to that on the Ti sample whereas the cells barely survived on the AZ31 or LDH coating. Furthermore, heparin was immobilized on LDH/PDA via a covalent bond (LDH/PDA/HEP). The corrosion resistance and long-term proliferation of HUVECs after the introduction of heparin were mildly decreased compared with the L/P sample, but were still greatly improved compared with AZ31, the LDH coating and the PDA coating. Furthermore, the LDH/PDA/HEP sample greatly improved the HUVEC migration rate compared with the LDH/PDA sample, and inhibited platelet adhesion which was intense on

Received 13th March 2018, Accepted 25th April 2018 DOI: 10.1039/c8bm00298c rsc.li/biomaterials-science

the LDH/PDA sample. Both LDH/PDA and LDH/PDA/HEP samples had a low hemolysis rate (2.52% and 0.65%, respectively) in vitro and eliminated the adverse biocompatible effects of the direct PDA coating on the AZ31 substrate in vivo. Our results suggest that the LDH/PDA composite coating with further heparinization is a promising method to modify the surface of Mg alloys by significantly improving corrosion resistance, endothelialization and hemocompatibility.

Introduction Coronary stents made from inert metals were introduced to clinical practice thirty years ago. Various improvements have been developed since then, such as drug-eluting technology. However, the problems of late restenosis and pulsatile restric-

a

Department of Cardiac Surgery, Zhongshan Hospital, Fudan University, Shanghai 200032, China. E-mail: [email protected] b State Key Laboratory of High Performance Ceramics and Superfine Microstructure, Shanghai Institute of Ceramics, Chinese Academy of Sciences, Shanghai 200050, China. E-mail: [email protected] c University of Chinese Academy of Sciences, Beijing 100049, China † Electronic supplementary information (ESI) available: Scanning maps of Mg, O, Al and C elements of all treated samples. The water contact angles of AZ31, LDH, PDA, LDH/PDA and LDH/PDA/HEP. Viability of HUVECs incubated for 3 days with different concentration extracts of AZ31, PDA, LDH, LDH/PDA and LDH/PDA/HEP. See DOI: 10.1039/c8bm00298c ‡ Co-first author.

This journal is © The Royal Society of Chemistry 2018

tion from permanent implants still persist. Considering the potential benefit of complete disappearance of implanted stents after the vessel-healing process, biodegradable polymerbased or metal-based stents have been developed since a decade ago and a number of clinical trials have been launched.1 Although various types of polymer-based stents being developed greatly outnumbered metal-based stents, the 2–3 year results of several high-quality clinical trials2,3 showed that polymer-based stents had higher rates of late luminal loss, device-oriented adverse events and late thrombosis compared with the metallic drug-eluting stents. Characteristics of polymers such as lower radial force, delayed resorption (>3 to 4 years) and excessive pro-inflammation contributed to the disappointing outcomes.4 The future of biodegradable stents may lie in the innovation of metal materials.5–7 Unlike polymers, magnesium (Mg) possesses superior physical properties such as low density (1.74 g cm−3), elastic modulus close to that of cortical bone (45 GPa vs. 5–23 GPa),

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favorable compressive strength (20–115 MPa) and tensile strength (90–190 MPa).8 Moreover, as the corrosion product, Mg2+ is involved in a series of life activities.9–11 However, Mgbased stents degraded rapidly in vivo,12 producing local accumulation of hydrogen bubbles and alkaline ions.13 Since the progression of negative vessel remolding lasted for more than 4 months after coronary angioplasty,14 biodegradable stents should provide adequate radial strength to support the vessel during this period. The fast degradation and the early loss of structural support contributed to the clinical failure of an early generation Mg alloy stent.15 Because of the excessive Mg ions and increased pH value, direct contact with Mg or Mg alloys exhibited minimal cell growth and no cell attachment in vitro.16–19 Our results also found that human umbilical vein endothelial cells (HUVECs) couldn’t survive on AZ31. Furthermore, platelets could adhere on the surface of Mg materials, which led to thrombosis and restenosis.20 Clinical pathology of drug-eluting stents has already demonstrated the associations of delayed endothelialization as well as in-stent thrombosis and restenosis.21 Therefore, it is essential for the Mg-based coronary stents to improve the anticorrosion and endothelialization. Surface modifications and bulk treatments are two general strategies to improve Mg corrosion resistance.20,22–25 Unlike the bulk treatments, surface modifications can build a thin coating on the surface of Mg, inducing behavioral alterations without changing the physical characteristics and chemical compositions of the Mg substrate.13 Surface modifications applied on Mg or Mg alloys include plasma electrolytic oxidation,26 phosphate treatment,27 alkaline heat treatment,28,29 anodization treatment,30 organic coating,31,32 etc. Not commensurate with the improvement of anti-corrosion by various coatings, the results of surface cytocompatibility were unsatisfied with short-term survival of seeded cells.33–35 Furthermore, one of the indispensable requirements of vascular stents is anticoagulation ability, which could be improved by surface modification.36,37 Though various coating methods for Mg have been reported, there were only a few research studies17,34,38–40 focusing on the multiple-functional composite coating. The present study endeavored to fabricate a composite coating on the AZ31 alloy with multiple-task performances (corrosion resistance, endothelialization and hemocompatibility) which are essential for coronary stents. Our previous work demonstrated that a Mg–Al layered double hydroxide (LDH) coating on Mg alloys showed enhanced corrosion resistance and biocompatibility in vivo and in vitro.13,17 Also, LDH coating didn’t produce the problem of gas pockets under the coating which could happen on polymer coatings.38 Nevertheless, ECs could not proliferate on the LDH coating. In the present study, an Mg–Al LDH/polydopamine (LDH/PDA) composite film was developed on the AZ31 alloy via hydrothermal treatment and the soaking method. The formation of a poly-dopamine (PDA) coating is ascribed to self-polymerization of dopamine, including oxidation of catechol to quinone, and then reaction with amines or catechols/quinines. The PDA coating showed improved cytocompatibility.41,42 Due to the existence of reactive functional

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groups (hydroxyl and amine), poly-dopamine is a desirable platform for further modification. Therefore, we further immobilized heparin on the surface of poly-dopamine via covalent bonding between amine/hydroxyl of poly-dopamine and carboxyl of heparin (LDH/PDA/HEP). Heparin is commonly used in clinic for its anticoagulant activity and has been shown to increase endothelial cell migration and blood compatibility.43,44 With the composite coating LDH/PDA to enhance corrosion resistance and cytocompatibility, and further heparinization to improve hemocompatibility and the migration of ECs, it is possible to produce a multiple-task coating of Mg-based implants. Herein, the corrosion resistance on coatings was investigated via electrochemical detection in vitro; the cytocompatibility to ECs was evaluated by culturing HUVECs on samples in vitro; the hemocompatibility was proved by the hemolysis rate and platelet adhesion tests in vitro; the biocompatibility was tested by subcutaneous implantation in vivo.

Experimental Specimen preparation Commercially purchased AZ31 alloy plates (10 mm × 10 mm × 2 mm) were used as substrates. The substrates were ground using 600 and 1000 silicon carbide papers and then ultrasonically cleaned in alcohol. The cleaned substrates were kept in alcohol to prevent oxidation and dried with an air blower before use. An Mg-Al LDH coating was synthesized on the AZ31 alloy via hydrothermal treatment according to our previous study.13 Briefly, AZ31 alloy plates were placed in a Teflon liner, and 50 mL aluminum nitrate solution (0.02 M, pH value was adjusted to 12.8 with NaOH) was gently poured into the Teflon liner. Then the Teflon liner was kept at 120 °C for 12 hours. After hydrothermal treatment, the samples were washed with a copious amount of deionized water. The obtained samples were denoted as LDH. To prepare a self-assembled poly-dopamine coating, the AZ31 alloy and LDH samples were subjected to 5 mg mL−1 dopamine hydrochloride (Aladdin) in 10 mM Tris-HCl ( pH 8.5) for 12 hours at 37 °C. The obtained samples were ultrasonically cleaned with ultrapure water to remove the loose poly-dopamine and were named PDA and LDH/PDA, respectively. To immobilize heparin (HEP, Aladdin) on the surface of the LDH/PDA sample, the sample was immersed in HEP solution (10 mg ml−1) for 2 hours at 37 °C. The obtained samples were named LDH/PDA/HEP. Coating characterization Scanning electron microscopy (SEM, Hitachi-S3800N, Hitachi, Japan) was used to observe samples’ surface morphology, and energy dispersive spectrometry (EDS, IXRF-550i, IXRF SYSTEMS, USA) was used to evaluate samples’ element composition. A digital camera was used to obtain the optical images of the samples. X-ray diffraction (XRD, JEM-2100F, JEOL Ltd,


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Tokyo, Japan) was performed to analyze the phase compositions of the samples.


Electrochemical corrosion evaluation An electrochemical analyzer (CH1760C, Shanghai, China) was utilized to investigate the corrosion behavior of all the samples. A three-electrode-cell was used with a saturated calomel electrode (SCE) as the reference electrode, a graphite rod as the counter electrode and the sample with an exposing area of 0.255 cm2 as the working electrode. Prior to the potentiodynamic polarization (PDP) and impedance spectroscopy (EIS) tests, the sample was stabilized in phosphate buffer saline (PBS) for 400 seconds. The PDP test was performed at a scanning rate of 10 mV s−1. The corresponding corrosion potential (Ecorr) and current density ( jcorr) were determined by Tafel extrapolation. The EIS test was performed with a 5 mV sinusoidal perturbing signal and the data were recorded from 100 kHz to 10 mHz. ZView software was used to analyze the impedance results. Cytoskeleton staining Human umbilical vein endothelial cells (HUVECs, ScienCell, USA) were used. Samples were sterilized via ultraviolet irradiation. HUVECs were seeded on AZ31, LDH, PDA, LDH/ PDA and LDH/PDA/HEP samples at a density of 5 × 104 cells per well. Samples were respectively taken out after 1, 4 and 24 hours of incubation, and then rinsed with PBS. For cytoskeleton staining, cells were fixed with 4% paraformaldehyde diluent, permeabilized with 0.1% (v/v) Triton X-100 (Amresco, USA) and blocked with 1 wt% bovine serum protein (BSA, Sigma, USA), respectively. Then, the cells were cultured with fluorescein isothiocyanate-phalloidin (FITC-phalloidin, SigmaAldrich) and 4′,6-diamidino-2-phenylindole (DAPI, SigmaAldrich) for 1 hour to stain F-actin and the nucleus, respectively. Finally, the samples were photographed using a confocal laser scanning microscope (CLSM, Leica SP8, Germany). Live-dead staining HUVECs were seeded (5 × 104 cells per well) on the samples for 3 days. Then, calcium-AM and propidium iodide (PI) were diluted in PBS with a final concentration of 2 and 5 μM, respectively, and 100 μL diluted solution was added to each sample. After being cultured for another 15 min, the cells were inspected by CLSM. Cell morphology and proliferation HUVECs were seeded (5 × 104 cells per well) on the AZ31 alloy, LDH, PDA, LDH/PDA and LDH/PDA/HEP samples. At the scheduled time (1, 3, 5 days), the samples were rinsed with PBS and fixed with 2.5% glutaraldehyde. After that, the samples were dehydrated in a grade ethanol series and hexamethyldisilizane ethanol solution series according to our previously published protocols.13 Then, the samples were sputtercoated with gold and observed by SEM. The proliferation rates of cells on LDH, LDH/PDA and LDH/ PDA/HEP were tested via the AlamarBlue assay. In detail, at the


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scheduled time (1, 3 and 5 days), samples were transferred to a new cell-culture plate. Then, both the new cell-culture plate and the old cell-culture plate were added with 0.5 mL AlamarBlue assay (diluted 10-folds with culture medium), and cultured for another 2 hours. Soon after that, 100 μL culture medium of each well was transferred to a black 96-well plate and measured using an enzyme-labeling instrument (BIO-TEK, ELX 800) with an absorption wavelength of 560 nm and a scattering wavelength of 590 nm. The fluorescence density of the measured culture medium was positive correlation with the cell proliferation rate. Cytocompatibility evaluation Live/dead cell staining was carried out to evaluate the biocompatibility of LDH, LDH/PDA, LDH/PDA/HEP and Ti samples. The live/dead cell staining kit was implemented as mentioned above after culturing for 10 days. Cell migration A scarification method was used to evaluate the migration of HUVECs. Cells were seeded on LDH, LDH/PDA, LDH/PDA/HEP and pure titanium (Ti) samples at a density of 20 × 104 cells per well. According to the result of cell adhesion, cells were spread on the surface after 4 h incubation. Therefore, after 4 hour incubation, culture media were removed and straight lines were drawn on the samples with a pipet tip. After that, 1 mL fresh culture medium was added to each well and cultured for another 8 hours. The cell staining process was the same as mentioned in the cell adhesion test. Finally, cell migration on different samples was observed through CLSM. Hemolysis rate Fresh human blood was obtained from healthy donors. Ethical approval was obtained from the research ethics committee of Zhongshan Hospital, Fudan University (Shanghai, China). Animal care and experimental protocols were conducted in accordance with the guidelines established by the Shanghai Medical Experimental Animal Care Commission. All donors were recruited in our study after informed consent. Physiological saline and distilled water were used as the negative control and positive control, respectively. Whole blood (0.8 mL) was diluted with physiological saline (1 mL). Prior to the test, samples were rinsed with PBS, and then immersed in 1.5 mL physiological saline for 30 minutes at 37 °C. Subsequently, each well was added with 30 μL diluted blood and kept for another 60 minutes. After that, the solution was centrifuged at 3000 rpm for 5 minutes. The absorbance of the supernatant at 545 nm was detected using an enzyme-labeling instrument. The hemolysis rate (HR) was calculated as follows: HR ¼

AS545 AN545 100% AP545 AN545

where AS545 represents the absorbance of the sample, AN545 represents the absorbance of the negative control and AP545 represents the absorbance of the positive control.

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Platelet adhesion Platelet-rich plasma (PRP) was obtained via centrifuging (15 minutes, 1500 rpm) fresh whole human blood. Each sample was added with 50 μL fresh PRP and incubated at 37 °C for 2 hours. Subsequently, the samples were rinsed with physiological saline (10 minutes, 3 times) and fixed with 2.5% glutaraldehyde. After being dehydrated in a grade ethanol series and hexamethyldisilizane ethanol solution series, the samples were sputter-coated with gold and observed by SEM.

Eight adult SD rats ( purchased from the Experimental Animal Centre of Zhongshan hospital), each weighing ∼200 g, were anesthetized with pentobarbital sodium (40 mg kg−1) by intraperitoneal injection. Five incisions were made on the back of each rat. One sample was implanted in the subcutaneous pocket through one incision. The samples on eight rats were harvested at 6 and 12 weeks after implantation, which provide four samples for each type of sample at each time point. After the capsule tissues containing the samples were resected, they were photographed with a digital camera and fixed in 10% neutral formaldehyde solution for 2 h. The tissues were dehydrated and embedded in paraffin. Histological cross-sections (∼5 μm) were stained with hematoxylin–eosin. Images were obtained using a bright-field microscope (DM6 M, Leica). The experiment was approved by the Animal Care and Use Committee of Zhongshan Hospital, Fudan University (Shanghai, China). Animal care and experimental protocols

Table 1

Statistical analysis Statistical analysis was performed using one-way ANOVA or two-way ANOVA. All of the data are expressed as the mean ± standard deviation (SD). A P-value