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Biodegradable Zwitterionic Polymer Coatings for Magnesium Alloy Stents Sang-Ho Ye, Yingqi Chen, Zhongwei Mao, Xinzhu Gu, Venkat Shankarraman, Yi Hong, Vesselin Shanov, and William R Wagner Langmuir, Just Accepted Manuscript • DOI: 10.1021/acs.langmuir.8b01623 • Publication Date (Web): 30 Jul 2018 Downloaded from http://pubs.acs.org on July 30, 2018
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Biodegradable Zwitterionic Polymer Coatings for Magnesium Alloy Stents
Sang-Ho Yea,b, Yingqi Chena, Zhongwei Maoa, Xinzhu Gua, Venkat Shankarramana, Yi Hongd, Vesselin Shanove and William R. Wagnera,b,c,d*
a
McGowan Institute for Regenerative Medicine, Departments of bSurgery, cBioengineering and d
Chemical and Petroleum Engineering University of Pittsburgh, Pittsburgh, PA 15219, USA
e
Department of Bioengineering, University of Texas at Arlington, TX 76019, USA.
e
College of Engineering and Applied Science, University of Cincinnati, Cincinnati, OH 45221,
USA.
*To whom correspondence should be addressed. William R. Wagner, Ph.D. McGowan Institute for Regenerative Medicine 450 Technology Dr., Suite 300 Pittsburgh, PA 15219, USA E-mail:
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ABSTRACT Degradable metallic stents, most commonly comprised of Mg-based alloys, are of interest as an alternative to traditional metallic stents for application in the cardiac and peripheral vasculature. Two major design challenges with such stents are control of the corrosion rate and the acute presentation of a non-thrombogenic surface to passing blood. In this study, several types of sulfobetaine (SB)-bearing biodegradable polyurethanes were developed and assessed as physical, chemical and combination-type coatings for a model degradable Mg alloy, AZ31. For physical coatings, poly(ester sulfobetaine)urethane ureas, PESBUUs were synthesized using variable monomers that allowed the incorporation of a varying extent of carboxyl groups. Introduction of the carboxyl groups was associated with faster polymer degradation time. Simple physical coating of PESBUUs reduced macro- and microscopic thrombogenic deposition together with good stability of the coating attachment compared to a control coating of polylactic-co-glycolic acid. For PESBUUs incorporating carboxyl groups (PESBUUs-COOH), these groups could be converted to siloxane groups (PESBUUs-Si), thus creating polymers that could be surface reacted with the oxidized or phytic acid treated AZ31 surface. Chemical (silanization) attachment of these polymers reduced underlying alloy corrosion rates, but following the salination reaction with physical coating most reduced corrosion rates and protected the surface better from the consequences of oxidation occurring under the coating, such as blistering. The application of a multi-layered coating approach using a sulfobetaine-based biodegradable elastomer thus offers options for degradable metallic stent design where thromboresistance is desired in combination with a means to control both polymeric coating degradation rates and underlying alloy corrosion rates.
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1. INTRODUCTION Degradable vascular stents have elicited substantial interest over the past twenty years and have progressed to clinical trials for both polymeric and metallic based designs1–3. Degradable metallic stents have an inherent strength advantage, requiring less material and thus requiring less of a footprint at the site of delivery and on the delivery system. Most commonly these stents have been made from Mg alloys, although more recently Zn-based designs have been proposed4,5. Like their non-degradable counterparts, these stents must present an adequately nonthrombogenic surface to passing blood, and the delivery of anti-proliferative agents from a coating may be desirable. Unlike non-degradable stents, the period of degradation from both a mass loss and mechanical property loss perspective becomes a major consideration. Coatings for degradable metallic stents have thus become of interest, since such a coating might be used to provide a non-thrombogenic surface for blood contact, provide a reservoir for controlled release of a pharmaceutical agent, and act as a protection against premature mass loss by inhibiting corrosion in the early implant period.
At the current clinical forefront of coated, degradable metallic stents is the Biotronik’s drugeluting absorbable metal scaffold (DREAMS)6 which is composed of a Mg alloy and coated with a biodegradable polymer (e.g. polylactic-co-glycolic acid (PLGA) or poly-L-lactide (PLLA)). The coating is used as a reservoir to deliver paclitaxel or sirolimus7 and has demonstrated improved results versus an earlier version of the stent which presented a bare metal surface to the blood7,8. Of note, the polymer employed for coating the stent is not known for high thromboresistance, nor is it particularly elastic, which would be of value in maintaining coating integrity as the stent is expanded.
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Based on the current clinical state-of-the-art, there thus appears to be a need for the design of a biomaterial approach specifically suited for application as a degradable metallic stent coating, rather than use of a commonly employed polyester that, although well characterized, appears to be lacking in many critical features. An improved coating material would include such desirable features as inherent thromboresistance, elasticity compatible with the maintenance of the coating upon stent expansion, and molecular binding to the metallic substrate to minimize early corrosion activity generating local areas of coating detachment (blistering). Since the ideal period over which the metallic stent needs to maintain mechanical viability may vary with application site and stent design, it would be attractive if the coating chemistry could readily be altered to control the rate at which the coating itself degraded.
As a well-recognized hemocompatible coating material for blood contacting devices, zwitterionic phosphorylcholine (PC) or sulfobetaine (SB)-bearing polymers have been broadly studied. Zwitterionic PC-bearing polymers coatings have been applied to commercial blood oxygenators9 and catheters10. A non-degradable metallic vascular stent has been coated by a drug eluting non-degradable PC-bearing polymer (Endeavor Zotarolimus-Eluting Coronary Stent developed by Medtronic and approved by the FDA) (2008). In considering degradable metallic substrates, PC and SB macromolecules have been attached to Mg alloy substrates and shown to resist acute thrombotic deposition and improve corrosion resistance11. However, the durability of the effects provided by a molecular coating on Mg alloys could be questioned, and such a molecular coating would not provide the option to serve as a drug release reservoir.
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Previously, we reported biodegradable elastomeric coatings on Mg alloy that showed improved corrosion resistance and coating stability under both static and dynamic degradation conditions12. Elastomeric behavior in stent coating materials is attractive to resist delamination or fracture during stent expansion. Also, we reported non-thrombogenic biodegradable elastomeric polyurethanes incorporating zwitterionic PC or SB in the polymer backbone for application as cardiovascular tissue engineering such as for blood vessel scaffolds13,14.
To specifically develop thromboresistant coatings for degradable metallic stents, in this report we have synthesized poly(ester sulfobetaine)urethane ureas (PESBUUs) using variable monomers that allowed the incorporation of a varying extent of carboxyl groups (PESBUUsCOOH) that should exhibit variable mechanical and degradation properties. Further, the COOH groups could be converted to other functional groups (e.g. siloxane groups to form PESBUUsSi), thus creating polymers that could be reacted with the oxidized or phytic acid treated Mg alloy (AZ31 alloy) surfaces. These surface-reactive zwitterionic polymers were also investigated in the design of a coating system combining chemical (i.e. primer) and physical techniques that could improve the early stage resistance of coating to corrosion, while still providing a thromboresistant surface to passing blood.
2. MATERIALS & METHODS 2.1. Materials N-Butyldiethanolamine (BEDA) and 1,3-propanesultone (PS), dimethylolpropionic acid (DMPA), glutaric acid (GA), 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), NHydroxysuccinimide (NHS), (3-aminopropyl)trimethoxysilane (APSi), 2,2,2-Trifluoroethanol 5 ACS Paragon Plus Environment
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(TFE) were purchased from Sigma-Aldrich and used without further purification. Polycaprolactone diol (PCL-diol, Mw =2000, Sigma-Aldrich) and sulfobetaine diol (SB-diol) synthesized from BEDA and PS (Ye et al., 2014) were dried in a vacuum oven at 60°C prior to synthesis. Diisocyanatobutane (BDI) and putrescine (Sigma-Aldrich) were purified by vacuum distillation. Stannous octoate (Sn(Oct)2) was dried using 4Å molecular sieves. Dimethyl sulfoxide (DMSO, anhydrous≥99.9%, Sigma), 1,1,1,3,3,3-hexafluroisopropanol (HFIP, Oakwood Inc.), and other chemicals used as received, except as noted above. Poly (D, L-lacticco-glycolic acid) (PLGA, LA:GA = 50:50, Mw=45,000) was purchased from Sigma-Aldrich. Mg alloy (AZ31, 3% Al, 1% Zn) plate was purchased from Goodfellow Corp. The AZ31 stent samples (3 mm dia, 250 µm struts thickness) were fabricated by a photolithography and chemical etching technique with a 250 µm thick AZ31 foil. The patterned and etched foil was rolled to form a cylinder and welded by laser15.
2.2 Synthesis of various SB-bearing biodegradable polyurethanes with/without carboxyl groups Biodegradable poly(ester sulfobetaine)urethane urea (PESBUU) elastomers which have SB groups in the polymer backbone were synthesized by mixing the SB-diol and PCL-diol during the two-step solvent polymerization technique as previously reported13. PE(SB)UU-0 and PESBUU-50 refer to elastomers used in this study that were synthesized using different molar ratios of PCL-diol: SB-diol, 100:0 or 50:50 respectively. Further, alternative SB-bearing poly(urethanes) with carboxyl groups (PESBUU-COOH) were also synthesized by adding DMPA into the mixed diols (SB-diol: PCL-diol: DMPA = 60:30:10). Then, BDI and putrescine (diols: BDI: putrescine =1:2:1 molar ratio) were used for the PESBUU-COOH polymer and
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synthesized as previously described13. Another PESBU-COOH which does not include a urea bond in the structure was also synthesized by using glutaric acid (GA) as the chain extender instead of putrescine after the reaction with diols and BDI (1:1 molar ratio). The polymer synthesis routes are summarized in Scheme 1. After synthesis, polymers were precipitated in ether and then washed with DI water and isopropanol for 3 d. The chemical structure of the synthesized polymers was confirmed by 1H-NMR and the degradation behavior was evaluated by weight loss after incubation in a lipase solution (100 U/mL) to better approximate in vivo degradation, as previously reported13.
2.3 Simple physical dip coating and silanization coating on AZ31 To compare the polymeric coatings and the thrombogenicity of various coated Mg alloy samples, each of the polymers was simply coated on cleaned and polished AZ31 samples without further pre-treatment of the Mg alloy samples. Cleaned AZ31 disk (8 mm dia.) or stent samples were dip-coated (3x) with the polymers (1wt% in HFIP) followed by solvent evaporation12. To achieve a simple silanization-based polymer coating, a siloxane functionalized polymer (PESBUU-Si) was prepared by reacting a PESBUU-COOH and APSi with EDC/NHS catalyst in TFE solution. Then the cleaned AZ31 samples were immersed in the mixed solution and stirred for 24 hr after adding triethylamine (TEA, Sigma) catalyst to achieve chemical bonding between the PESBUU-Si and Mg(OH)2 via surface silanization11. Then the PESBUU-Si coated sample was simply washed with TFE and dried in vacuum. The PESBUU-Si was used as a priming layer 16
.
2.4 Phytic acid pretreatment of AZ31 prior to silanization (combination coating)
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To add corrosion resistance and surface hydroxyl groups on AZ31 samples, the surface was also pre-treated with phytic acid (PA, Shanghai Macklin Biochemical Co., Ltd) following an alkaline heat treatment (3M NaOH solution, 60 oC for 3hr)16. This was followed by a silanization reaction with PESBUU-Si. The PA coated AZ31sample (AZ31-PA) was placed in a PESBUU-Si solution (1 wt%) and further stirred for 24 hr after adding TEA as a catalyst. After the reaction, the sample was simply washed with TFE and dried in vacuum (AZ31-PA-PESBUU-Si, Chemical). Then, an additional top coating layer was also applied on the sample by polymer dip coating (x1) to complete the combination coating (AZ31-PA-PESBUU-Si-PESBU-COOH, Combination).
2.5 Thrombotic deposition analysis after ovine blood contact Whole ovine blood was collected by jugular venipuncture13. NIH guidelines for the care and use of laboratory animals were observed, and all animal procedures were approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Pittsburgh. In vitro thrombotic deposition on the sample was evaluated by employing a simple rocking test13,16 with fresh ovine blood (heparin 3.0 U/mL for acute studies or sodium citrate 10.6 mM for incubations lasting 4 hr). The surfaces were observed by scanning electron microscopy (SEM; JSM-6330F, JEOL USA, Inc., Peabody, MA). The thrombotic coverage on the surfaces was evaluated using Image-J software (NIH) with selected representative electron micrographs11.
2.6 Rat Vascular Smooth Muscle Cell (rSMC) cell growth The attachment and growth of rSMCs was evaluated on the polymer coated and uncoated AZ31 plates as previously reported12. Briefly, samples were sterilized under 30 min UV irradiation, then immersed in a well (24-well cell culture plate) loaded with culture medium (1 mL
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Dulbecco’s Modified Eagle’s Medium supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin solution). Cells (rSMCs) were seeded at 104 cells per well. After 3 h incubation to allow cell attachment, the specimens were then transferred into new wells and cultured for 7 days with culture medium exchanged every 3 days. Tissue culture polystyrene (TCPS) cut to the same dimension (10 x 5x 0.25 mm) served as a positive control group. The metabolic activity of rSMCs was measured using a MTS assay kit (Promega CellTiter 96 Cell Proliferation Assay) (n=4). To qualitatively verify the results of the MTS assay, cells were also observed under fluorescence microscopy after live/dead staining with a live/dead viability/cytotoxicity kit (Invitrogen Inc.), and images were taken using fluorescence microscopy (Eclipse Ti, Nikon).
2.7 Coating stability study (Adhesive behavior of coating on Mg substrate) The coated AZ31 surfaces were observed by SEM after ovine blood contact, but platelet poor plasma (PPP) solution with sodium azide (0.1%, as a preservative) was also utilized to investigate the coating stability over extended periods. To further investigate the adhesive behavior of coatings on Mg substrates, a scratch test17,18 was utilized to compare the simple polymeric coating and chemically bonded coatings. A grid of parallel cuts with 1 mm intervals were scratched using a sharp needle to leave a scissure on the coating. The scratched samples were then immersed in PPP solution at 37 ̊C and rocked (the PPP solution was refreshed every day). After three days of immersion, the samples were washed with DI water and dried in a vacuum chamber. The surface morphology, especially the interfaces between the polymer layer (scratched) and Mg substrate were observed using the SEM.
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2.8 Electrochemical corrosion measurement The corrosion properties were also evaluated by an electrochemical corrosion analysis16. Potentiodynamic polarization (PDP) curves and electrochemical impedance spectroscopy (EIS) were recorded using an electrochemical workstation (IM6, Zahner, Germany). A conventional three-electrode system was implemented, which involves a reference electrode (saturated calomel electrode, SCE), counter electrode (platinum sheet, 1.5 cm × 1.5 cm), and working electrode (test sample). For the working electrode, the backside of the samples was connected electrically through a copper wire and then was sealed with silicone rubber to expose the research surface (ca. 1 cm2 area). The sample connected with copper wire was immersed into phosphate buffered saline (PBS) electrolyte. Prior to potentiodynamic polarization test, the samples were immersed into PBS solution 30 min to stabilize the surface states, and then the PDP curves were scanned from -2.0 to -1.0 VSCE at a scanning rate of 1 mV/s. The corrosion potential Ecorr and current density Icorr were determined by the Tafel method through linear extrapolation of the cathodic polarization zone. Electrochemical impedance spectroscopy (EIS) was performed in a Faraday cage using the same electrochemical workstation. After the samples were stabilized in PBS for 30 min, the EIS spectra were recorded from 200 kHz to 0.01 Hz with a 10 mV (peak to peak) sinusoidal perturbing signal at open circuit potential. All electrochemical corrosion experiments were repeated on at least four different samples for statistical analyses.
All results are presented as mean ± standard deviation (SD). Data were analyzed by one-way ANOVA followed by a post hoc Newman-Keuls testing. Significant differences were considered to exist at p
