New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets

Acta Biomaterialia xxx (2013) xxx–xxx

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New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets Raila Busch a,⇑, Anne Strohbach a, Stefanie Rethfeldt a, Simon Walz a, Mathias Busch a, Svea Petersen b, Stephan Felix a, Katrin Sternberg b a b

University of Greifswald, Clinic for Internal Medicine B, Sauerbruchstraße, D-17475 Greifswald, Germany University of Rostock, Institute for Biomedical Engineering, Friedrich-Barnewitz-Strasse 4, D-18119 Rostock, Germany

a r t i c l e

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Article history: Received 7 May 2013 Received in revised form 2 September 2013 Accepted 14 October 2013 Available online xxxx Keywords: Stents Biodegradable polymers Endothelial cells Smooth muscle cells Platelets

a b s t r a c t Despite the development of new coronary stent technologies, in-stent restenosis and stent thrombosis are still clinically relevant. Interactions of blood and tissue cells with the implanted material may represent an important cause of these side effects. We hypothesize material-dependent interaction of blood and tissue cells. The aim of this study is accordingly to investigate the impact of vascular endothelial cells, smooth muscle cells and platelets with various biodegradable polymers to identify a stent coating or platform material that demonstrates excellent endothelial-cell-supportive and non-thrombogenic properties. Human umbilical venous endothelial cells, human coronary arterial endothelial cells and human coronary arterial smooth muscle cells were cultivated on the surfaces of two established biostable polymers used for drug-eluting stents, namely poly(ethylene-co-vinylacetate) (PEVA) and poly(butyl methacrylate) (PBMA). We compared these polymers to new biodegradable polyesters poly(L-lactide) (PLLA), poly(3hydroxybutyrate) (P(3HB)), poly(4-hydroxybutyrate) (P(4HB)) and a polymeric blend of PLLA/P(4HB) in a ratio of 78/22% (w/w). Biocompatibility tests were performed under static and dynamic conditions. Measurement of cell proliferation, viability, glycocalix width, eNOS and PECAM-1 mRNA expression revealed strong material dependency among the six polymer samples investigated. Only the polymeric blend of PLLA/P(4HB) achieved excellent endothelial markers of biocompatibility. Data show that PLLA and P(4HB) tend to a more thrombotic response, whereas the polymer blend is characterized by a lower thrombotic potential. These data demonstrate material-dependent endothelialization, smooth muscle cell growth and thrombogenicity. Although polymers such as PEVA and PBMA are already commonly used for vascular implants, they did not sufficiently meet the criteria for biocompatibility. The investigated biodegradable polymeric blend PLLA/P(4HB) evidently represents a promising material for vascular stents and stent coatings. Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Although the use of stents has represented a revolutionary advance in the treatment of cardiovascular disease, in-stent restenosis (ISR) and stent thrombosis (ST) remain major limitations of coronary intervention. ISR occurs in arteries treated with bare metal stents in 15-30% of percutaneous coronary interventions (PCIs) [1–3]. The pathophysiological process of ISR is due to uncontrolled neointimal hyperplasia and tissue proliferation in response to vessel injury that occurs during balloon deployment of the stent. The degree of neointima formation can be considerably reduced if

⇑ Corresponding author. Tel.: +49 03834/8680500; fax: +49 03834/8680502. E-mail address: [email protected] (R. Busch).

re-endothelialization occurs quickly and completely after intervention. Furthermore, patient factors such as diabetes and small vessel diameter [4,5], as well as stent surface characteristics, shape properties and elastic recoil, are determinants in the formation of ISR [6]. However, the occurrence of ISR has been reduced to <5% since the introduction of drug-eluting stents (DESs) into clinical routine [2,3]. A great number of available DESs are coated with a permanent polymeric carrier loaded with an anti-proliferative and antiinflammatory drug that is eluted via diffusion-controlled processes [6,7]. The first DES was the sirolimus-eluting CypherÒ stent (Johnson & Johnson, USA), approved by the Food and Drug Administration in 2003, which was based on a blend of two permanent polymers: poly(ethylene-co-vinylacetate) (PEVA) and poly(butyl methacrylate) (PBMA) in ratio of 50/50% (w/w). The coating

1742-7061/$ - see front matter Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.actbio.2013.10.015

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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contains 1.4 lg mm 2 of the immunosuppressant sirolimus and a thin layer of pure PBMA as top coat [8]. Despite improvement in clinical outcome involving major adverse cardiac events [9,10], patients treated with a DES can still suffer from late ST [11–13]. Although a very rare complication, ST can lead to severe myocardial infarction or death. Furthermore, cases of local hypersensitivity were observed in connection with DES implantations [14]. Due to recent technical progress in stent development, second-generation DESs produce much less ST and have been shown to be superior over the first generation in many studies for safety and efficacy [15–17]. The ongoing CHOICE trial is now investigating whether there are differences in efficacy and safety among second-generation DES after long-term follow-up [18]. Recent studies suggest that the more rapid the re-endothelialization of the stented area, the lower the risk of inflammation and subsequent thrombus formation [19]. This raises the concern in DES usage that—in addition to non-specific drug action on the vascular tissue—the polymeric coating materials themselves may inhibit the growth of endothelial cells, leading to disturbed vascular healing. In this context, it was stated that the stent platform, the drug and the polymeric drug carrier are all targets for DES improvement [20]. In vivo endothelial cells are responsible for various anti-platelet, anticoagulant, fibrinolytic and growth-regulatory properties. These range from release of prostacyclin, inactivation of thrombin by surface-bound thrombomodulin, synthesis of plasminogen activator and secretion of growth factors that can inhibit smooth muscle cell growth [19]. Accordingly, a polymer material that enables complete re-endothelialization of the upper surface layer, and can in turn promote normal smooth muscle cell growth underneath, could prove of great value for new stent technologies. Moreover, a material that also has the capacity to simultaneously promote endothelial cell attachment could generate a surface that is actively non-thrombogenic and non-inflammatory. Recent activities in second-generation DES development are targeting more biocompatible and biodegradable polymer carriers—or even fully absorbable stent platforms that degrade during or after drug release and would not cause long-term foreign-body reactions. The polymers principally used are polylactides (PLAs) [21–24] as well as their copolymers, which are fully metabolized to carbon dioxide and water (Krebs cycle) [25]. The first clinically investigated absorbable DES platform is the balloon-expandable BVS stent (Abbott Vascular, USA), made of poly(L-lactide) (PLLA). The coating of the BVS stent consists of an amorphous poly(D,L-lactide) (PDLLA) matrix incorporated with everolimus [26]. Published one-year optical coherence tomography data demonstrated that the everolimus BVS stent (31 lesions) induces a neointimal response similar to that of the metallic everolimus DES (19 lesions) (Xience, Abbott Vascular, USA) [27]. Separately, biodegradable stents of poly(3hydroxybutyrate) [28] and of blends of poly(4-hydroxybutyrate) (P(4HB)) and PLLA have also been developed [29] and tested preclinically. In this context, the aim of this in vitro study was to evaluate biocompatibilities of various stent surfaces. We consequently investigated interactions among human endothelial cells, smooth muscle cells (SMCs) and platelets with the biodegradable stent coating and platform materials PLLA, P(3HB), P(4HB), and a blend of PLLA and P(4HB) in the ratio of 78/22% (w/w). We furthermore studied the interactions between the various cell types with the permanent polymers PEVA and PBMA contained in the coating of the CypherÒ stent. All experiments were referred to Thermanox™, which is known for its high biocompatibility and is therefore used as control surface. In addition, we conducted static investigations and dynamic flow experiments using an in vitro perfusion model to analyze cell–material interactions under the preferred physiological conditions. Flow conditions of the venous as well as the

arterial system (shear stress from 1.5 to 20 dyn cm lated by means of this cell perfusion chamber.

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) were simu-

2. Methods 2.1. Polymer film preparation The following polymer base materials were used during film preparation: PEVA (Mw = 45,000 g mol 1, Arkema GmbH, Düsseldorf, Germany), PBMA (Mw = 337,000 g mol 1, Sigma–Aldrich, Taufkirchen, Germany), PLLA (ResomerÒ L214, Mw = 720,000 g mol 1, Boehringer Ingelheim Pharma, Ingelheim, Germany), P(4HB) (TephaFLEXÒ, Mw = 170,000 g mol 1, Tepha Inc., Lexington, MA, USA) and P(3HB) (Mw = 760.000 g mol 1, Hans Knöll Institut, Jena, Germany). Whereas we employed a pouring process for PEVA, PBMA, P(4HB), and the polymeric blend composed of 78 wt.% PLLA and 22 wt.% P(4HB), we applied manual dip-coating for PLLA and P(3HB). For film pouring, 1 g of the corresponding polymer base material was dissolved in 25 mL chloroform (Mallinckrodt Baker, Griesheim, Germany) and poured into a glass Petri dish (9 cm diameter). Due to the high molecular weight of PLLA, a lower concentration of 0.52 g PLLA and 0.15 g P(4HB) in 25 ml chloroform was used for blend preparation. After pouring, the solvent was allowed to evaporate until formation of a film 100 lm thick, as measured by a thickness gauge (2109 Mitutoyo, Mitutoyo Europe GmbH, Neuss, Germany). Subsequently, the films were cut out of the Petri dish and washed in distilled water for 5 days and dried for a further 7 days in a vacuum drier at 40 °C and 40 mbar. For dip-coating, concentrations of polymer base materials were 1.68 g for PLLA and 4.00 g for P(3HB) in 100 ml chloroform. The dipping process was repeated with intermediate drying of 10 min until films 100 lm thick were obtained on cylindrical substrates. After cutting the polymer off the substrate, washing and final drying were performed as described above for the poured polymer films. Prior to in vitro testing, all films were sterilized by a common ethylene oxide sterilization process and afterwards stamped out under sterile conditions into disks of 25 and 5 mm diameter for flow chamber and static experiments, respectively. 2.2. Cell culture Human umbilical vein endothelial cells (HUVECs) were obtained from collagenase type II (Biochrom, Berlin, Germany) digested umbilical cords and cultivated as described previously [30]. Primary cultures of human coronary artery endothelial cells (HCAECs) and human coronary artery smooth muscle cells (HCASMCs) were purchased from Provitro GmbH, Berlin, Germany. Cells were obtained at passage 2 and cultured in endothelial cell growth medium (Provitro GmbH) or smooth muscle cell growth medium (Provitro GmbH) supplemented with 10% fetal bovine serum (FBS, Invitrogen, Carlsbad, CA, USA). Cells were cultured for 4 days before treatment or exposure to shear stress. 2.3. Flow chamber experiments For flow chamber experiments, cells at passages 3–6 were seeded into 6-well cell culture plates (Greiner-bio one) at a density of 2  105 cells on 25 mm diameter Thermanox™ coverslips (Nunc, Langenselbold, Germany) or on polymer films, and were cultured at 37 °C in a humidified mixture of 5% CO2 and 95% air. HUVECs, HCAECs and HCASMCs were cultured to 80% confluence, placed in a parallel-plate flow chamber (FCS, Provitro GmbH) and perfused for up to 3 h at 37 °C with 2% FBS supplemented medium

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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or platelet-rich plasma (PRP). Cells were subjected to static conditions and to defined low and high levels of pulsatile laminar shear stress (1.5 and 20 dyn cm 2, respectively). Control cells not exposed to shear stress were also cultured in 2% FBS-supplemented medium for the same length of time as the shear-stress-treated cells. After shear-stress exposure, cells and control cells were stained or detached with Accutase™ (eBioscience, San Diego, CA) for further characterization. 2.4. Cell proliferation The relative proliferation rates were quantified by BrdU Cell Proliferation ELISA (Roche, Mannheim, Germany). Endothelial and smooth muscle cells were seeded at a density of 6,000 cells on Thermanox™ coverslips or polymer films with a diameter of 5 mm, and were cultured for 6 days in black 96-well microplates with flat, clear bottoms. According to the manufacturer’s instructions, 100 lM BrdU was added to each well and incubated for 24 h at 37 °C, 5% CO2 and 95% humidity. Cell proliferation was determined using an Infinite M200 multimode microplate reader (Tecan, Männedorf, Switzerland). 2.5. DAPI and PECAM-1 staining Representative DAPI images of HUVECs adhered to control and polymer surfaces were recorded under a magnification of 20. HUVECs were cultured in 6-well plates on glass coverslips and polymer films (P(3HB), PLLA, P(4HB) and PLLA/P(4HB) (78/22%, w/w)) for 4 days under static conditions. Cells were stained for platelet endothelial cell adhesion molecule 1 (PECAM-1) with primary antibody rabbit anti-CD31 (1:200, Santa Cruz Biotechnology, Santa Cruz, CA) and secondary antibody anti-rabbit IgG (Alexa FluorÒ 555) (1:1000, Cell Signaling Technology, Danvers, MA, USA). Nuclei were stained for DAPI. Pictures displayed are representative for at least three separate experiments. 2.6. Cell viability Cell viability was assessed under static and dynamic conditions. Under static conditions, cells were seeded onto 5 mm diameter Thermanox™ coverslips or polymer films at a density of 6,000 cells per 96-well and matched to untreated controls. Under dynamic conditions, cells were seeded on Thermanox™ coverslips or polymer films and placed in the flow chamber system as described above. After exposure to dynamic flow conditions, cells were detached with Accutase™, seeded into a 96-well plate and matched to untreated controls of exactly the same cell number. AlamarBlueÒ (Biosource, Camarillo, CA) was added to each well as 10% of the sample volume, and cells were incubated for 16 h. The resulting fluorescence was read on a Wallac 1420 VICTOR2™ plate reader (Perkin Elmer, Waltham, MA) using an excitation wavelength of 570 nm and an emission wavelength at 585 nm.

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Serial optical sections were acquired from the bottom to the top of the sample in defined steps of 0.29 lm. Cell nuclei were stained for DAPI. 2.8. RNA Isolation, cDNA synthesis and real-time PCR Total RNA was isolated using an RNeasy Mini Kit (Qiagen). Reverse transcription was carried out with 200 ng of total RNA by random hexamer primers and TaqMan reverse transcription reagents (Applied Biosystems, Carlsbad, CA). The synthesized single-strand cDNA was used to quantify GAPDH, NOS3 and PECAM-1 expression levels with a 7900HT Fast Real-Time PCR System (Applied Biosystems) using the following TaqMan gene expression assays (Applied Biosystems): GAPDH (4352339E), NOS3 (Hs00167166_m1) and PECAM-1 (Hs00169777_m1). Relative gene expression data were normalized to the amount of the housekeeping gene GAPDH in the same cDNA preparation, by using the 2 – DDCt method. 2.9. Preparation of platelets Human platelets were obtained from the Institute for Immunology and Transfusion Medicine, University of Greifswald, Germany, in a polypropylene bag containing acid citrate dextrose A (an anticoagulant solution, 1/6 volume of blood). The upper PRP layer was obtained by centrifuging platelet concentrates at 150g for 15 min. PPP was obtained after centrifuging the concentrate at 2000g for 15 min. We added 2.5 mM CaCl2 and 1 mM MgCl2 to the PRP and incubated for 30 min at 37 °C. PRP was then used for static and shear-stress experiments. 2.10. Scanning electron microscopy (SEM) and evaluation of platelet adhesion Glass coverslips and polymer films were immersed in PBS for 30 min in 6-well plates to equilibrate the surface and were then incubated with PRP (2  108 cells ml 1) for 2 h at 37 °C. Coverslips were washed three times with PBS to remove non-adherent platelets. After a fixation step (1 h in 2.5% glutaraldehyde in PBS), samples were treated with 1% osmium tetroxide in PBS for 1 h at room temperature, dehydrated in a graded series of acetone, and then critical-point dried. Finally, samples were mounted on aluminum stubs, sputtered with gold–palladium, and examined in an EVO LS10 scanning electron microscope (Zeiss, Oberkochen, Germany). Surface evaluation was acquired using a high-vacuum technique at 500 and 5000 magnification. Platelet adhesion on five polymer surfaces and glass as control was analyzed using ImageJ software (NIH, Bethesda, MD). Particles with a diameter of 10 lm and above were considered aggregates. An average of three representative images (500) was obtained per polymer and used for quantitative analysis of platelets and platelet aggregates.

2.7. Confocal laser scanning microscopy (CLSM) 2.11. ELISA (sP-selectin) Primary endothelial cells (HUVECs) were cultured in 6-well plates on glass coverslips and polymer films (P(3HB) and on PLLA/P(4HB) (78/22%, w/w). Cells were exposed to different levels of laminar shear stress for 3 h at 37 °C. The corresponding controls were maintained under static conditions. Lectin binding (lectin from Triticum vulgaris WGA-FITC or lectin from Lycopersicon esculentum LEA-FITC) was observed with a confocal laser scanning microscope (TCS SP5 X, Leica Microsystems GmbH, Wetzlar, Germany). All CLSM images (20448  2048 pixels) were acquired using an oil immersion objective lens (Leica, 63  1.4 oil). To observe lectin–FITC binding, an argon laser (k = 488 nm) was used.

Human platelets were isolated as described above. Before shear-stress exposure, pump tubing was rinsed with PBS for 5 min and BSA (4 mg ml 1) for 2 min. Control (Thermanox™ and glass) and polymer surfaces were exposed to pulsatile laminar shear stress and perfused with PRP (2  108 cells ml 1) for 3 h at 37 °C. After perfusion, PRP was collected and centrifuged at 1000g for 15 min. To determine platelet activation by polymer contact, sP-selectin enzyme-linked immunosorbent assay (ELISA) (R&D Systems, Abingdon, UK) was performed as defined by the manufacturer.

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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2.12. Data collection and statistical analysis

3. Results

All experiments were repeated at least three times, and representative experiments are shown. For statistical analysis, IBM SPSS Statistics 19 software was used. Alterations in gene and protein expression after shear-stress exposure were compared to untreated controls using a non-parametric Kruskal–Wallis test. Statistical differences between controls and shear-stresstreated cells or between treatments were analyzed using a non-parametric Mann–Whitney U-test. Unless otherwise stated, P < 0.05 was used to indicate statistical significance for all data analyses.

3.1. Proliferation of human endothelial cells and smooth muscle cells on various polymer materials under static conditions Fig. 1A, B shows the relative proliferation of endothelial cells (HUVECs and HCAECs) and smooth muscle cells (HCASMCs) on the investigated polymer materials under static conditions. Thermanox™ is used as control surface and data are set at 100%. The proliferation of HUVECs and HCAECs is similar on the various polymers, as demonstrated in Fig. 1A. Proliferation of HUVECs and HCAECs on P(3HB) is greater than on the Thermanox™ control (Fig. 1A, ⁄P < 0.05), while it is almost doubled on PLLA/P(4HB)

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Fig. 1. Proliferation of human endothelial cells and smooth muscle cells on various coating materials under static conditions. Primary endothelial cells (A, HUVECs j and HCAECs j) and smooth muscle cells (B, HCASMCs) were cultured in 96-well plates on Thermanox™ (control) and polymers PEVA, PBMA, PLLA, P(4HB), P(3HB) and PLLA/ P(4HB) blend for 6 days at a density of 6,000 cells per well. Endothelial and smooth muscle cells were maintained under static conditions. (A,B) Bars show mean ± SD of at least three separate experiments after BrdU incubation for 24 h, normalized to a Thermanox™ control. ⁄P < 0.05; ⁄⁄P < 0.01 and ⁄⁄⁄P < 0.001. (C) Dot blot shows distribution of proliferation of HCASMCs with HUVECs (left) and HCAECs (right).

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Fig. 2. Human endothelial cells grown on various coating materials under static conditions. Representative fluorescence images of HUVECs adhered to control and polymer surfaces (20). HUVECs were cultured in 6-well plates on glass (control) and on polymers P(3HB), PLLA, P(4HB) and PLLA/P(4HB) blend for 4 days under static conditions. Cells were stained for endothelial cell marker PECAM-1 (red) and nuclei were stained for DAPI (blue).

(⁄P < 0.05). Proliferation of HUVECs and HCAECs on PEVA, PBMA, PLLA and P(4HB) is less than on Thermanox™. In addition to the BrdU test, DAPI staining of cell nuclei (20) was performed to confirm the proliferation data. Fig. 2 shows examples of HUVECs on the control surface, P(3HB), PLLA, P(4HB) and PLLA/P(4HB). Simultaneously, immunostaining of the endothelial cell marker PECAM-1 displayed cell junctions and spreading of the endothelial cells on different surfaces. HUVECs on control surface and PLLA/P(4HB) exhibited a complete monolayer on the material. In contrast, PECAM-1 staining of HUVECs on P(3HB) showed an irregular distribution, and cells lay in clusters on the material. Furthermore, HUVECs do not form a monolayer on PLLA and P(4HB) when used separately, and PEVA or PBMA surfaces induce only a few insular cells (data not shown). The pattern of HCASMC proliferation on polymers resembles the pattern of endothelial cells (Fig. 1B). Proliferation of HCASMCs on P3(HB) and PLLA/P(4HB) is enhanced compared to the control (⁄P < 0.05) and reduced on P(4HB) and PEVA (⁄P < 0.05). Fig. 1C displays a box-plot diagram of the distribution of proliferation of HCASMC and endothelial cells (data from Fig. 1A, B). Under static conditions, there is no material that simultaneously shows more proliferation of endothelial cells and less proliferation of HCASMCs than the control (top left quadrant).

3.2. Viability of human endothelial cells and smooth muscle cells on various polymer materials under dynamic conditions The following results refer only to HUVECs, since they are commonly used in the literature as a stable model of endothelial cells. Data obtained for HCAECs are for the most part equivalent to those for HUVECs but are not shown here. Fig. 3A demonstrates the relative cell viability (data of Thermanox™ are set as 100%) of HUVECs and HCASMCs on Thermanox™, PEVA, PBMA, PLLA, P(4HB),

P(3HB), and PLLA/P(4HB) under static conditions. The highest viability of HUVECs is achieved on PLLA (⁄⁄P < 0.01), whereas HUVECs on all other polymers show less viability compared to the static Thermanox™ control (⁄⁄P < 0.01). Fig. 3B demonstrates the relative cell viability for selected polymers under flow conditions in comparison to the corresponding static polymer control (#). There is no difference for PLLA compared to the corresponding static PLLA control. The highest viability of HUVECs is achieved on the PLLA/P4(HB) under high shear stress, which correlates to flow conditions in a coronary artery (#P < 0.05). The relative cell viability of HCASMCs on Thermanox™ under shear stress is less than under static conditions (Fig. 3C, ⁄ P < 0.05). Cell viability of HCASMC on P(4HB) under flow conditions is enhanced in comparison to its corresponding static control (Fig. 3C).

3.3. Glycocalyx width of human endothelial cells on different polymer materials under static and dynamic conditions Fig. 4 shows examples of glycocalyx width determinations of HUVEC monolayers on Thermanox™, P(4HB), P(3HB) and PLLA/ P(4HB). Laser scanning microscopy images (Fig. 4A) support these measurements (Fig. 4B and C). Under static conditions glycocalyx width on control and PLLA/P(4HB) is between 2 and 3 lm, which corresponds to physiological data. Under shear-stress conditions, the glycocalyx on the control surface decreases (⁄P < 0.05; significance data on the corresponding static control are not shown in Fig. 3C). In contrast, HUVECs grown on PLLA/P(4HB) showed a decreased glycocalyx width under high shear stress only (⁄P < 0.01). Glycocalyx width on P(3HB) is augmented under static conditions (⁄P < 0.05) and grows further under low and high shear stress (⁄P < 0.05). Glycocalyx on P(3HB) morphologically appears clumpy and thickened (see also Fig. 2). It was not possible to measure

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Fig. 3. Viability of human endothelial cells and smooth muscle cells on various coating materials under flow conditions. (A) Primary endothelial cells (HUVECs) and smooth muscle cells (HCASMCs) were cultured in 6-well plates on the surface material for 4 days (control: Thermanox™). Cells were maintained under static conditions. Bars show mean ± SE for at least three separate experiments after Alamar BlueÒ incubation for 16 h normalized to a static control. ⁄P < 0.05. (B, C) Cells were exposed to levels of laminar shear stress (j 1.5 dyn cm 2 and j 20 dyn cm 2) for 3 h at 37 °C. After shear-stress exposure, cells were detached and seeded into 96-well plates. Bars show mean ± SE for at least three separate experiments after Alamar BlueÒ incubation for 16 h, normalized to the static Thermanox™ control (⁄P < 0.05) and the corresponding static control (#P < 0.05).

glycocalyx width on PBMA and PEVA under static or under shearstress conditions because of poor endothelialization. Measurement and staining of glycocalyx on PLLA was also not feasible because of the bowing of the material. 3.4. Expression of the shear-stress-dependent endothelial cell markers eNOS and PECAM-1 on various polymer materials under flow conditions Endothelial cell function analyses were performed by measuring the stress-dependent cell markers endothelial nitric oxide synthase (eNOS) and PECAM-1 (CD-31) as positive indicators of successful formation of cell–cell junctions and interactions, under static and shear-stress conditions. Fig. 5A demonstrates relative mRNA expression of eNOS on control, PLLA and PLLA/P(4HB) under static and shear-stress conditions. Measurement of the eNOS mRNA of P(4HB) as a component of the blend, PEVA, PBMA and P(3HB) was not possible because of the low cell number and low mRNA amount on these materials. As already described [29], data of the control confirm an increase in eNOS expression under shear stress (⁄⁄P < 0.01). Interestingly, eNOS expression under static and shear-stress conditions is dependent on the material (⁄P < 0.05). eNOS expression on PLLA and PLLA/P(4HB) is higher than on the control and ascends farther on PLLA/P(4HB), with increasing shear stress, to approximately 4 times its static value (#P < 0.05).

PECAM-1 expression of endothelial cells also differs depending on the surface material. PECAM-1 expression of HUVECs on PLLA/ P(4HB) is enhanced under static conditions compared to the Thermanox™ control (Fig. 5B; HCAEC data not shown). Data recently described [30] that show that PECAM-1 increases under shear stress have been confirmed for the control data (⁄⁄P < 0.01). Furthermore, under shear-stress conditions, PECAM-1 expression of endothelial cells on PLLA and PLLA/P(4HB) has a trend to increased levels, which demonstrates the capability of the cells to create cell junctions and spread under physiological conditions. 3.5. Platelet adhesion and aggregation on various polymer materials Platelet adhesion and aggregation, as a marker of the prothrombotic potential of polymers, was visualized by exposing platelets to all six polymers for 30 min (see above). Fig. 6A shows representative SEM images of adhered and aggregated platelets on Thermanox™ and polymer surfaces (500 and 5000 magnification). Quantitative analysis of platelets and platelet aggregates confirms the SEM findings (Fig. 6B). On the control surface, platelets demonstrate normal activation states. The highest number of platelets was seen on P(4HB) and the highest platelet aggregation, on PLLA. Despite the fact that platelet adhesion and aggregation on PEVA is relatively small, platelets appear morphologically conspicuous without any formation of pseudopodia on this polymer. The

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Fig. 4. Glycocalyx width of human endothelial cells on various coating materials under static and dynamic conditions. Primary endothelial cells were cultured in 6-well plates on glass (control) and polymer coverslips (P(3HB), P(4HB) and PLLA/P(4HB) blend) for 4 days. Cells were exposed to various levels of laminar shear stress for 3 h at 37 °C. Corresponding controls were maintained under static conditions. (A) Confocal fluorescence images of HUVECs stained with lectin WGA-FITC in various focal planes. Serial optical sections were acquired from the bottom to the top of the sample in defined steps of 0.29 lm. Green fluorescence on the cell surface indicates lectin binding to the glycocalyx. (B, C) Glycocalyx width of HUVECs grown on polymer surfaces and exposed to static conditions and shear stress. Bars show mean ± SD of at least three separate experiments after lectin staining, normalized to the corresponding Thermanox™ control (⁄⁄⁄P = 0.001).

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Fig. 5. Expression of shear-stress-dependent endothelial cell markers eNOS and PECAM-1 on PLLA and PLLA/P(4HB) blend under static and flow conditions. HUVECs were cultured in 6-well plates on Thermanox™ (control) and polymer coverslips (PLLA and PLLA/P(4HB) blend) for 4 days. Cells were exposed to various levels of laminar shear stress (j 1.5 dyn cm 2 and j 20 dyn cm 2) for 3 h at 37 °C. Corresponding controls were maintained under static conditions. After shear-stress exposure, cells were detached and prepared for qRT-PCR. Bars show mean ± SD of at least three separate experiments normalized to a static Thermanox™ control (⁄P < 0.05) and the corresponding static control (#P < 0.05 and ##P < 0.01).

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Fig. 6. Platelet adhesion on various coating materials under static conditions. (A) Representative SEM images showing adhered platelets on a control surface (glass) and various polymer surfaces (500 and 5000). Polymers were incubated with PRP (2  108 ml 1) for 2 h at 37 °C. Platelets show polymer-dependent adhesion patterns (upper panel) and activation states (lower panel). (B) Bars show median of the number of adhered platelets and aggregates on various polymer coatings of at least three independent experiments.

smallest adhered platelet number and the least platelet aggregation was demonstrated on PLLA/P(4HB). The polymer blend may thus contribute to more non-thrombotic surface conditions.

3.6. Platelet activation by means of sP-selectin on various polymer materials under static and flow conditions sP-selectin measurement as a marker of platelet activation was performed after static incubation of PRP on various polymers (Fig. 7A) and under flow conditions (Fig. 7B). There was no difference in relative sP-selectin concentration between the polymers and the Thermanox™ control (set as 100%). After shear-stress exposure, sP-selectin concentration increases on PEVA and P(3HB) compared to their corresponding static controls (#P < 0.05). Platelets on PEVA are more activated after low and high shear stress than the other polymers: sP-selectin concentration after low shear stress is almost doubled (#P < 0.05) in comparison to its corresponding static control. There was no change in platelet activation between flow and static conditions on P(4HB), PLLA and PLLA/P(4HB).

4. Discussion Although the first generation of DESs has revolutionized the treatment of coronary artery disease in terms of significant reduction of ISR rates, there is increasing evidence that applied permanent polymer coatings such as PEVA and PBMA could be responsible for adverse effects such as ST and residual restenosis [31–34]. In the context of developing second-generation DESs, biodegradable polymers, copolymers and polymeric blends based on polyglycolide, PLA, P(4HB) and P(3HB) are being investigated [35–37]. However, tissue response to the foreign materials in terms of vessel healing, inflammation and biocompatibility has not yet been described in detail. In this context, the present study was designed to compare endothelialization, smooth muscle cell proliferation and thrombogenicity on permanent PEVA and PBMA in comparison to biodegradable PLLA, P(3HB), P(4HB), and a customized blend of PLLA/ P(4HB), by using a flow chamber model with laminar shear-stress application. These materials are intended as stent coatings or platform materials for the stent structure. We found material-dependent platelet adhesion and activation as well as endothelial cell and smooth muscle cell growth. Additionally, a qualitative difference of endothelial cell function was measured by eNOS and PE-

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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A

B

Fig. 7. Platelet activation on various coating materials under flow conditions. (A) Polymer surfaces were incubated with PRP (2  108 ml 1) for 3 h at 37 °C under static conditions. (B) Polymer surfaces were incubated with PRP (2  108 ml 2) for 3 h at 37 °C under dynamic conditions. After incubation, PRP was collected and centrifuged, followed by determination of sP-selectin concentration by ELISA. Bars show mean ± SD of at least four separate experiments, normalized to the corresponding static control (#P < 0.05 and ##P < 0.01).

CAM-1 mRNA expression. In this context, the bioabsorbable blend PLLA/P(4HB) evidently demonstrates remarkable endothelial cellsupportive properties that are well suited for clinically justified stent deployment. Most studies analyzing the impact of material on blood or tissue involve experiments under static conditions, e.g. cell incubation with the material of interest. In contrast, the flow chamber model used in this study generates mechanical forces in vitro that are similar to those produced by physiological hemodynamic forces in vivo. A major mechanical force—wall shear stress—is relevant in the mimicking of physiological conditions for blood and tissue cells, in particular when an artificial device is implanted [38]. To simulate the diameter and blood flow of a typical coronary artery, we applied a shear stress of 20 dyn cm 2 and a temperature of 37 °C. Low shear stress of 1.5 dyn cm 2 corresponds to venous conditions. Before collecting data for this study, flow chamber model validation measurements, including cell viability and endothelial NOS under shear stress, were performed (data not shown). The viability of endothelial cells treated with shear stress was always >85%. Endothelial regrowth is an essential factor after stenting, because a continuous and complete endothelial layer prevents

inflammatory and thrombotic events on the vessel wall. Besides antiplatelet and fibrinolytic properties, a healthy endothelium is responsible for growth regulation by secreting growth factors that can inhibit smooth muscle cell proliferation [39]. Moreover, Finn et al. reported that the most powerful histological predictor of ST is endothelial coverage [40]. They observed that heterogeneity of healing is a common finding in DESs with evidence of late ST, and they demonstrate the importance of incomplete healing of the stented segment in the pathophysiology of late ST. On the basis of our observations, proliferation of endothelial cells on PBMA, PEVA, PLLA and P(4HB) is less than on the control surface, whereas endothelial cell proliferation on PLLA/P(4HB) is almost doubled compared to the control. Moreover, endothelial cells grown on the PLLA/P(4HB) surface exhibit an organized cytoskeleton, enhanced eNOS expression and a normally developed glycocalyx. Endothelial cells on PLLA/P(4HB) adhere as a strong continuous monolayer and show a distinct structure organization. In contrast, PEVA or PBMA surfaces induce few insular cells with poor viability. Although HUVEC proliferation on P(3HB) is high, cells do not compose a monolayer but grow in disordered clusters. This observation could explain why HUVEC viability under static and low shearstress conditions is still relatively high (60% higher than on

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Thermanox™; data not shown), but under high shear stress is 50% less than with the control. In healthy vessels the endothelial glycocalyx functions as a vasculoprotective layer over the endothelium [41,42]. In our study, lectin labeling of the glycocalyx of HUVECs and subsequent CLSM revealed a surface layer thickness of 2.5 ± 0.5 lm. Measurement of the glycocalyx width of HUVECs and HCAECs on control and PLLA/P(4HB) under static and flow conditions was always within physiological ranges [43]. Unfortunately, it was not possible to measure glycocalyx width on PBMA and PEVA because of poor endothelialization on these materials. Measurement of glycocalyx width of HUVECs subject to shear stress on P(3HB) yielded dimensions 3–4 times those of normal control. Glycocalyx on P(3HB) morphologically appeared clumpy and thickened, presumably because of the cluster-like cell disposition. Evidence is emerging that damage to the glycocalyx plays a pivotal role in several vascular pathologies such as diabetes and atherosclerosis [44]. When the glycocalyx is injured, disrupted or modified, its vasculoprotective properties are lost, as has been shown in recent experimental settings [44]. Coronary stent implantation leads to a period of decreased flow and to a disrupted endothelium with microvascular dysfunction. Since the glycocalyx is injured after coronary intervention, endothelial cells do not have an efficient barrier. These data suggest that the endothelial glycocalyx may function as a marker for the endothelialization capability of the surface material. Here, we were able to show for the first time that eNOS expression of endothelial cells is dependent on the material on which they are cultivated. eNOS, as a marker of endothelial cell function, represents the metabolic status of the cell. Previous studies have shown that eNOS is reduced with endothelial dysfunction, e.g. hypertension and arteriosclerosis [45]. In addition, Kawashima et al. demonstrated that eNOS overexpression inhibits lesion formation in a mouse model of vascular remodeling [46]. Endothelial recovery in stented coronary arteries strongly depends on the stent surface material. A polymer material on which eNOS production is high under shear-stress conditions is more qualified than a material that rarely triggers cell functionality. In the present study, we also examined mRNA expression of the endothelial antigen PECAM-1, a transmembrane glycoprotein that functions as a cell–cell adhesion molecule. Once a cell–cell contact is constituted, PECAM-1 accumulates on contact sites and, via homophilic extracellular domains, establishes interactions with neighboring cells [47]. PECAM-1 plays a key role in platelet-leukocyte/endothelial-cell interaction and in forming and maintaining the contact-inhibited state in endothelial cells. Endothelial cells form cell junctions during processes such as cell growth and migration, displaying PECAM-1 immunostaining at endothelial layers [48,49]. Joner et al. documented a generally poor PECAM-1 expression of sirolimus- and paclitaxel-eluting stent struts after 14 and 28 days, suggesting an inhibition of endothelial cell migration and proliferation, endothelial injury and/or increased cell turnover [51]. In this study, PECAM-1 mRNA expression of endothelial cells was different on diverse polymers. In comparison to the static control, endothelial cells on the polymer blend PLLA/ P(4HB) showed elevated PECAM-1 expression under static conditions. It has recently been shown that PECAM-1 expression is upregulated under shear stress [38,50], which we confirmed here by data for the control surface. Data of this study demonstrate furthermore that endothelial expression of PECAM-1 is dependent on the surface material. Compared to PLLA, endothelial cells on PLLA/ P(4HB) present high PECAM-1 expression, which suggests that PLLA/P(4HB) has the ability to establish cell–cell junctions and interactions for a healthy endothelial layer. Although this observation was made under in vitro conditions, a pulsatile shear stress of 20 dyn cm 2 simulates in vivo flow conditions in a coronary artery. Shear stress is detected by luminal mechanosensors, and signals

are transmitted through the cytoskeleton to the basal or junctional surface where certain integrins or a mechanosensory complex consisting of PECAM-1 and Flk-1 are activated, respectively, and initiate a downstream signaling cascade [38]. Arterial injury-like PCI with stent implantation leads to changes in shear-stress conditions on the vessel wall [38]. Moreover, an incomplete endothelium that is not able to regrow after denudation cannot transmit mechanical forces into the cell, which leads to endothelial dysfunction and remodeling. The composition and texture of the subjacent polymer material could also effect changes in mechanical tensions of the endothelial cells and therefore influence the process of mechanotransduction. Consequentially, a material that allows adequate and complete endothelial recovery could support the physiological process of mechanotransduction. Another aspect concerning cell–material interaction involves inflammatory reactions on the vessel wall that could influence the biocompatibility of the stent surface [51,52]. We in fact observed material-dependent monocyte rolling and adhesion (data not shown). In this context, Faigle et al. stated that, especially under hypoxia, ATP release and metabolism at the interface between polymorphonuclear and vascular endothelial cells may influence the inflammatory reaction [53]. Presently used DESs that release antiproliferative drugs reduce smooth muscle cell growth but also enhance ST by suppressing reendothelialization. Improvements in PCI procedures include the use of second-generation DESs, together with new coating technologies, bioabsorbable stents and non-drug-based stent coatings. Particular emphasis will be placed on the concept that endothelial regeneration could be pursued as well as a reduction of smooth muscle cell proliferation to allow stable successful revascularization after DES deployment [39]. Data from this study emphasize the fact that smooth muscle cell proliferation is dependent on the material on which they were grown. In our study we demonstrated that PBMA, PEVA and P(4HB) are polymers that do not aggravate smooth muscle cell proliferation, whereas P(3HB), PLLA and PLLA/P(4HB) accelerate growth. Unfortunately, we did not identify a material that enhances endothelial proliferation while at the same time reducing smooth muscle cell proliferation. Nevertheless, smooth muscle cell proliferation tests were performed only under static conditions, since incubation with BrdU is not possible under flow. Interestingly, viability tests after high shearstress exposure show that PLLA/P(4HB) is a material with good viability properties for both cell types. Interaction of platelets with the material is critical in the process of vascular healing since deposition of platelets on polymeric surfaces can complicate endothelialization and thus facilitate thrombotic events. Following activation, platelets release their pro-coagulant a-granules that contain P-selectin, which is subsequently expressed on the surface of the activated platelets and released into the plasma (soluble P-selectin), where it can be used as a differential marker for activation. Unfortunately, it is not clearly understood whether platelets are indeed activated while they cross coronary lesions. Yong et al. reported direct evidence of up-regulation of platelet P-selectin, platelet–monocyte aggregate formation and monocyte CD11b expression in human coronaries, in proportion to shear stress and stenosis severity [54]. However, Yin et al. exposed platelets to high shear stress of 60 dyn cm 2 for 0.1 s every 90 s, a protocol that did not induce significant changes in platelet activation or in platelet thrombogenicity potential [55]. These results suggest that—even though the presence of relative severe stenosis (at 60 dyn cm 2) can affect local flow conditions and can greatly increase local shear stress (from 10 to 60 dyn cm 2)—this does not necessarily lead to platelet activation or thrombin generation. These observations partly correspond with our results. We found material-dependent platelet activation under shear stress. A shear stress of 20 dyn cm 2 led to sP-selectin

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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activation on Thermanox™, PEVA, PBMA and PLLA. Platelets which were in contact with P(3HB), P(4HB) and PLLA/P(4HB) did not exhibit activation compared to their static control. Interestingly, the high sP-selectin expression of PEVA does not correlate with the amount of adhered or aggregated platelets. Blit et al. observed a high degree of sP-selectin expression in platelets that were in direct contact with polyethylene (PE) [56]. PEVA is a copolymer of ethylene and vinylacetate. The question of whether structural PE properties are the reason for platelet activation must be reserved for further investigation. Moreover, the low sP-selectin expression after contact with PLLA/P(4HB) correlates well with the lower amount of adhered and aggregated platelets, and furthermore with a significant reduction of collagen-induced aggregation measured by the method of Born [57] (data not shown here). We assume that the component P(4HB), or its monomer c-hydroxybutyric acid (4-HB), may be responsible for the non-thrombogenic qualities. Recently, Franconi et al. showed that 4-HB as human endogenous metabolite reduces platelet aggregability in a dose-dependent way [58]. The underlying question of why P(4HB) has these properties only in combination with PLLA, but not as single material, should be investigated in further studies on the basis of structural analysis. PLLAs of various molecular weights are frequently used as stent polymers. Lincoff et al. evaluated the inflammatory response to Wiktor stents coated with low (80 kDa) and high (321 kDa) molecular weight PLLA in pig coronary arteries [59]. In the group with low molecular weight PLLA, severe acute and chronic signs of inflammation were recognized with variable destruction of the vessel wall architecture. In contrast, in the group with high-molecular-weight PLLA, there was no evidence of acute or chronic inflammation, and the neointima was similar to the control group. Because our study involves only short-term experiments (<1 week), no influence of molecular weight on biocompatibility could have been expected from this point of view. Using PLLA designs, Zidar et al. revealed that stents have shown promising in vitro mechanical properties with nearly complete degradation and with minimal inflammatory response after 9 months. A further study assessed the technical feasibility and impact on coronary stenosis of a paclitaxel-eluting PDLLA stent in porcine coronary arteries [35]. Coronary stenosis after implantation of paclitaxel-eluting stents was inhibited compared to drug-free PDLLA stents and metal stents after 3 weeks and 3 months. Although early endothelialization became apparent, local inflammatory response to the polylactide was observed. This observation is also in accordance with Bünger et al., who found that a sirolimus-eluting PLLA stent demonstrated less sign of inflammation in porcine arteries than did the drug-free PLLA stent [21]. In this study, venous and arterial endothelial cells cultivated on PLLA with an average molecular weight of 650 kDa showed the same proliferation and cell viability as on the control surface, especially under shear stress. With the purpose of optimizing mechanical properties of the stent material, P(4HB) was added to PLLA in a weight ratio of 78/22% [60]. The blend PLLA/P(4HB) is less rigid than PLLA alone, degrades faster and also demonstrates improvement of biocompatibility as shown in this study. The two non-erodible polymers PEVA and PBMA are coating constituents of the first DES, the sirolimus-eluting CypherÒ stent (Johnson & Johnson, USA). Delayed endothelialization in combination with incomplete stent apposition to the vessel wall subjects patients to a higher and longer risk for ST [12,61,62]. Furthermore, Hofma et al. observed endothelial dysfunction 6 months after sirolimus-eluting stent implantation [63]. The authors suggested that the anti-proliferative activity of sirolimus may result in prolonged healing response, with concomitant delayed recovery of endothelial function.

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Our PEVA/PBMA data showed a very poor endothelial cell proliferation rate, particularly for HCAECs. Accordingly, the viability of endothelial cells on PEVA and PBMA was low (20–30% compared to the control surface; data not shown here). Measuring eNOS release on PEVA and PBMA was not possible because of poor proliferation, as mentioned above. Although platelet adhesion and aggregation on PEVA and PBMA were less than with most of the other polymers, platelets appeared morphologically conspicuous without signs of normal activation states such as spreading and formation of pseudopodia. In addition, sP-selectin concentration was significantly increased, especially after high shear-stress exposure. Hence, not only the immunosuppressive effect of drugs such as sirolimus and biolimus—as well as incomplete apposition—but also inappropriate conditions for endothelial cells to grow on the PEVA/ PBMA evidently lead to endothelial dysfunction, attachment of inflammatory cells and ST. In this respect, it would be very interesting to study the endothelialization process of a surface coating consisting of an immunosuppressive drug and a polymer that promotes proliferation of endothelial cells. It would, moreover, be logical to subsequently perform vasomotion studies to evaluate the timeframe of endothelial functional recovery. In light of these data, it would be reasonable to compare differences regarding crystallinity, surface energy and surface morphology of the favored polymeric blend PLLA/P(4HB) compared to PLLA, P(4HB) and other PLLA/P(4HB) blend compositions with regard to thrombogenicity and endothelial and smooth muscle cell growth. For further studies with PLLA/P(4HB), it will also be of great interest to investigate co-cultures of involved blood and tissue cells after material implantation in an adjusted perfusion model. In this regard, it would represent an appreciable advance to find a biomaterial for which intrinsic endothelial cell growth is superior to smooth muscle cell proliferation and hyperplasia, without application of an immunosuppressant drug. 5. Conclusion Although the first generation of DESs has revolutionized the treatment of coronary artery disease in terms of significant reduction of ISR rates, there is increasing evidence that applied permanent polymer coatings such as PEVA and PBMA could be responsible for adverse effects such as delayed healing, late ST, local hypersensitivity reaction and residual ISR. We therefore investigated biodegradable polymers and blends thereof. However, the tissue response to these materials in terms of vessel healing, inflammation and biocompatibility have not yet been described in detail. In this context, the present study was designed to compare permanent and biodegradable polymers with the use of human endothelial and smooth muscle cells under laminar shear stress in an in vitro flow chamber model. Our data disclose material-dependent endothelialization, smooth muscle cell growth and thrombogenicity. Although polymers such as PEVA and PBMA are already commonly used for vascular implants, they did not meet sufficient criteria for biocompatibility. The investigated polymeric blend PLLA/P(4HB) might well represent a promising material for vascular stents and stent coatings. The extent of smooth muscle cell growth and spreading should be tested, however, in an in vivo model before clinical application. Acknowledgements We thank Roswitha Dressler (University of Greifswald, Clinic for Internal Medicine B) for excellent assistance in RTq-PCR and Western blot procedure. We are likewise obliged to Ulrike Jehmlich (University of Greifswald, ZIK HIKE: Innovation Center—Humoral

Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Immune Responses in Cardiovascular Disorders) for providing assistance with confocal microscopy. Our appreciation also goes to Dr. Rabea Schlüter (University of Greifswald, Institute for Microbiology) for SEM exposures, Andrea Rohde (University of Rostock, Institute for Biomedical Engineering) for polymer film preparation and Provitro GmbH (Berlin, Germany) for technical support. Additionally, we gratefully acknowledge Dr. David Martin (Tepha Inc., Lexington, MA, USA) for supply of the biodegradable biomaterial P(4HB), as well as for his helpful notes and suggestions. Financial support by the Bundesministerium für Bildung und Forschung (BMBF, German Federal Ministry of Research and Technology) within REMEDIS ‘‘Höhere Lebensqualität durch neuartige Mikroimplantate’’ (FKZ: 03IS2081, see http://remedis.med.uni-rostock.de for details) is gratefully acknowledged.

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Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015

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Please cite this article in press as: Busch R et al. New stent surface materials: The impact of polymer-dependent interactions of human endothelial cells, smooth muscle cells, and platelets. Acta Biomater (2013), http://dx.doi.org/10.1016/j.actbio.2013.10.015