In vitro and in vivo evaluation of a dextran-graft-polybutylmethacrylate copolymer coated on CoCr metallic stent

Introduction

Percutaneous transluminal coronary angioplasty followed by a stent implantation is the first technique of the revascularization. Most of these stents are manufactured using cobalt chromium coated with iridium oxide or diamond-like carbon layers to prevent the formation of an oxide layer on their surface that is fragile and mechanically unstable under flow conditions.¹ Despite the development of new coronary stent technologies, interactions of blood and cells with the implanted devices remain a serious cause of side effects. The major complications of their implantation are restenosis and late stent thrombosis (LST), which raises serious questions about their long-term safety. In-stent restenosis results from the proliferation of smooth muscle cells (SMCs).² Late thrombus formation mechanism has not been fully explained yet, nevertheless, LST is thought to be associated with premature cessation of antiplatelet therapy in patients, stent malapposition, late re-endothelialization of the stent struts and the nature of polymer coating.^3-6 In 45% of cases, they lead to patient death.

In this context, three main approaches were developed to design safer drug-eluting stent (DES).⁷ The first one is polymer-free coating.^8-10 Stent surfaces are micro-porous or contain cavities where drugs are loaded. Nevertheless, the release control on these stents is difficult and the velocity is often too high.^8,11,12 The second one consists of a biodegradable polymer coating that allows the drug-delivery during the degradation of the polymer.^10,13 In this case, the polymer and its degradations products have to be biocompatible and not trigger any inflammatory reactions. Many biodegradable coatings were investigated.^14,15 Alternatively, biodegradable stents were also developed.^11,16,17 Nevertheless, side effects such as inflammatory reactions related to their degradation products have been reported.¹⁵ Clinical studies revealed that these bioresorbable stents led to an increased risk of stent thrombosis in comparison with DES. As a matter of fact, metallic stents appear to be the most adapted device to prevent coronary artery closure. The third approach is therefore based on a non-degradable polymer that protects the stent from corrosion. In vitro metallic stent backbone was shown to release degradation products leading to a possible allergic reaction or sensitization.¹⁸ The polymer must be biocompatible and should promote endothelial coverage to prevent side effects in contact with blood.⁵ Recent studies suggest that the faster the re-endothelialisation of the stented area is, the lower the risks of inflammation and thrombus formation are.¹⁹ The drug action and the polymeric coating materials themselves may inhibit the growth of endothelial cells, leading to disturbed vascular healing.²⁰ In this context, it was stated that the stent platform, the drug and the polymeric drug carrier are all targets for DES improvement.¹⁹

Polymethacrylate polymers are widely used on medical devices for their mechanical properties and particularly the poly(n-butylmethacrylate) (PBMA) on several DES.²¹ Nonetheless, it is associated with another polymer because if the polymer is used on its own, its elasticity is too low to resist to stent deformation.^21,22 Previous works have demonstrated that a copolymer of dextran, a natural polysaccharide, along with PBMA (Dex-PBMA ) exhibits elastic properties suitable for stent coating.²² This combination promotes in vitro endothelial cell growth and inhibits the SMCs proliferation compared to PBMA alone.²³ Besides, dextran is non-toxic and has already been approved by the US Food and Drug Administration for its clinical use for example as plasma volume expander.²⁴

Therefore this study was undertaken to evaluate the haemocompatibility and biocompatibility of Dex-PBMA coated on glass coverslips and CoCr discs or stents, in vitro on coagulation, platelets and endothelial cells. Then more studies were performed ex vivo and in vivo in preclinical models of restenosis and intimal hyperplasia.

Material and Methods

Results

Dex-PBMA characterization

The synthesis leads to a homogeneous white powder of dextran-graft-polybutylmethacrylate copolymer (Dex-PBMA ) soluble in a mixture of THF and water (92:8 v/v). The coatings of Dex-PBMA on a glass coverslip, CoCr disc and CoCr stent or moulding resulted in a transparent film, homogeneous and without cracks. Dex-PBMA deep-coating of stent covered each strut independently with very few webbing. The stent surface modifications were studied using SEM and contact-angle analysis. As shown Fig. 1A, this coating withstands to the crimping on the balloon and the deployment. After one week of incubation at 37°C and under coronary-like flow conditions, Dex-PBMA coating remained unaltered and showed a uniform polymer surface, niether pealing nor delamination was observed.

The coating wettability was evaluated by measuring water contact angles (Fig. 1B). CoCr is less hydrophobic (θ=88.7° ± 2.2) than Dex-PBMA with or without Tacrolimus (respectively θ=98.5° ± 0.4 and 99° ± 2.9, P = 0.0078).

Fig. 1

Dex-PBMA morphology after stent coating. (A) SEM micrographs of dextran grafted poly(butyl methacrylate) (Dex-PBMA ) coated stent. Stents were expanded, dip-coated, crimped on a balloon and expanded again. They spend 8 days in water either in static or under flow conditions, shear rate 1000 s^-1. (B) Water contact angles on cobalt chrome (CoCr), Dex-PBMA or Dex-PBMA with tacrolimus (Dex-PBMA _TAC) discs. Values are mean ± SEM of 10 determinations ** P<0.01.

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Fig. 1. Dex-PBMA morphology after stent coating. (A) SEM micrographs of dextran grafted poly(butyl methacrylate) (Dex-PBMA ) coated stent. Stents were expanded, dip-coated, crimped on a balloon and expanded again. They spend 8 days in water either in static or under flow conditions, shear rate 1000 s^-1. (B) Water contact angles on cobalt chrome (CoCr), Dex-PBMA or Dex-PBMA with tacrolimus (Dex-PBMA _TAC) discs. Values are mean ± SEM of 10 determinations ** P<0.01.

Bacterial adhesion

As shown Fig. 2, Staphylococcus aureus was lower on Dex-PBMA than on CoCr (ratio 11, P < 0.005) and control collagen (ratio 4.5, P < 0.05). In contrast, more bacteria adhered on CoCr than on control (ratio 4.7, P < 0.005).

Fig. 2

Bacterial adhesion on Dex-PBMA . Dex-PBMA film, CoCr disc and Collagen (Coll) were incubated 1h incubation at 37°C with a suspension of 108 UFC/mL of Staphylococcus aureus. After washing the remaining bacteria were cultivated on agar gel overnight and counted (n=3). Values are mean ± SEM of three determinations *P<0.05; ** P<0.01.

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Fig. 2. Bacterial adhesion on Dex-PBMA . Dex-PBMA film, CoCr disc and Collagen (Coll) were incubated 1h incubation at 37°C with a suspension of 108 UFC/mL of Staphylococcus aureus. After washing the remaining bacteria were cultivated on agar gel overnight and counted (n=3). Values are mean ± SEM of three determinations *P<0.05; ** P<0.01.

Adhesion of platelets and thrombin generation on Dex-PBMA

In static condition, platelet adhered less on Dex-PBMA (Fig. 3A) than on the positive control collagen (ratio 2.7, P < 0.0001) or on CoCr (ratio 1.39, P < 0,01). The SEM images showed less platelet activation on the copolymer than on the bare metal. Non-activated platelets are disk-shaped. Once activated, they first take a spherical shape as observed on Dex-PBMA disk (Fig. 3A) before spreading and giving out extensions such as filopodia that can be observed on CoCr surface.³² In dynamic conditions, no platelet adhered on Dex-PBMA (Fig. 3B). Dex-PBMA coating does not promote platelet adhesion and activation in our experimental conditions.

Thrombin generation was performed on Dex-PBMA and CoCr. In PPP, the velocity index of thrombin generation was 1.9 fold lower on Dex-PBMA disc than on CoCr disc (Fig. 3C, P < 0.01) in presence of Tissue Factor. There were no differences between the two surfaces for thrombin generation when the assay was performed in PPP without Tissue Factor (Fig. 3C) or in PRP with or without Tissue Factor (data not shown).

Fig. 3

Human platelet adhesion and thrombin generation onto Dex-PBMA . (A) Left panel: Platelet adhesion in static condition on Dex-PBMA or CoCr discs and bovine serum albumin (BSA, negative control). Quantification of adherent platelets after 2h incubation at room temperature. Data are expressed as a percentage of adherent platelets on collagen (positive control). Values are mean ± SEM of three determinations. **P<0.01. Right panel: Environmental scanning electron micrographs of adherent platelets on Dex-PBMA and CoCr. Scale bar 20 μm, higher magnification scale bar 2 μm. (B) Up: Dynamic adhesion of dioc-6 labeled platelets on Dex-PBMA and collagen coated glass-coverslips submitted to physiological flow (shear rate 1500 s^-1). Micrographs of adherent platelets after 0, 30, 60 and 120 s. Down: Quantification of the adhesion within times on Dex-PBMA and collagen. (C) Thrombin generation on Dex-PBMA and CoCr discs. Left panel: Representative curve obtained after a thrombin generation experiment. Parameters of thrombin generation are shown: Lagtime (LT in min), time to peak (tPeak in min) and Peak (nM). Right panel: Velocity Indexes (VI=Peak/(tPeak-LT) for thrombin generation in PPP in the absence or presence of TF (1 pM) on Dex-PBMA and CoCr. Data are expressed as a percentage of the results on well plate (WP). Values are mean ± SEM of six determinations. **P<0.01.

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Fig. 3. Human platelet adhesion and thrombin generation onto Dex-PBMA . (A) Left panel: Platelet adhesion in static condition on Dex-PBMA or CoCr discs and bovine serum albumin (BSA, negative control). Quantification of adherent platelets after 2h incubation at room temperature. Data are expressed as a percentage of adherent platelets on collagen (positive control). Values are mean ± SEM of three determinations. **P<0.01. Right panel: Environmental scanning electron micrographs of adherent platelets on Dex-PBMA and CoCr. Scale bar 20 μm, higher magnification scale bar 2 μm. (B) Up: Dynamic adhesion of dioc-6 labeled platelets on Dex-PBMA and collagen coated glass-coverslips submitted to physiological flow (shear rate 1500 s^-1). Micrographs of adherent platelets after 0, 30, 60 and 120 s. Down: Quantification of the adhesion within times on Dex-PBMA and collagen. (C) Thrombin generation on Dex-PBMA and CoCr discs. Left panel: Representative curve obtained after a thrombin generation experiment. Parameters of thrombin generation are shown: Lagtime (LT in min), time to peak (tPeak in min) and Peak (nM). Right panel: Velocity Indexes (VI=Peak/(tPeak-LT) for thrombin generation in PPP in the absence or presence of TF (1 pM) on Dex-PBMA and CoCr. Data are expressed as a percentage of the results on well plate (WP). Values are mean ± SEM of six determinations. **P<0.01.

Cell adhesion and proliferation onto Dex-PBMA

ECFCs and MSCs adhered less on Dex-PBMA disc than on CoCr disc (Fig. 4). Non-significant differences were obtained for intermediate adhesion (20 min incubation time) whereas for focal adhesion (2 hours incubation time), 2.4 less ECFCs (Fig. 4A, P < 0.0005) and 1.6 less MSCs (Fig. 4B, P < 0.05) adhered on Dex-PBMA compared to CoCr discs. The same trends were observed for cell proliferation (Fig. 4C and Fig. 4D). The proliferation of ECFCs and MSCs was greater on CoCr disc than on Dex-PBMA . In both cellular type and surfaces, the proportion of cells compared to control coating of gelatin decrease after 4 days. Thus, CoCr or Dex-PBMA surfaces cause cells to grow more slowly than on control gelatin wells.

Fig. 4

Circulating cells adhesion and proliferation on Dex-PBMA . (A-B) Human endothelial colony-forming cells (ECFCs) (A) and mesenchymal stem cells (MSCs) (B) intermediate (20 min) and focal (2 h) adhesion on Dex-PBMA (black bar) and CoCr (white bar) discs. (C-D) ECFCs (C) and MSCs (D) proliferations after 72h (day 3) and 96h (day 4) of culture. Data are expressed as a percentage of the positive control: gelatin. Values are mean ± SEM of 3 determinations. ***P < 0.001, **P < 0.01.

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Fig. 4. Circulating cells adhesion and proliferation on Dex-PBMA . (A-B) Human endothelial colony-forming cells (ECFCs) (A) and mesenchymal stem cells (MSCs) (B) intermediate (20 min) and focal (2 h) adhesion on Dex-PBMA (black bar) and CoCr (white bar) discs. (C-D) ECFCs (C) and MSCs (D) proliferations after 72h (day 3) and 96h (day 4) of culture. Data are expressed as a percentage of the positive control: gelatin. Values are mean ± SEM of 3 determinations. ***P < 0.001, **P < 0.01.

Release of tacrolimus from Dex-PBMA

Fig. 5 shows the cumulative TAC release profile from ex vivo stented arteries. An extreme burst release was observed corresponding to 5 ng/mL (50%) after 3 hours, followed by a steady state reaching 10 ng/mL over 4 days.

Fig. 5

Ex vivo cumulative Tacrolimus release profile from rat abdominal aorta stented with Dex-PBMA _TAC stent for a period of 2 hours in perfusion medium at 37°C under flow conditions (shear rate 1000 s^-1).

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Fig. 5. Ex vivo cumulative Tacrolimus release profile from rat abdominal aorta stented with Dex-PBMA _TAC stent for a period of 2 hours in perfusion medium at 37°C under flow conditions (shear rate 1000 s^-1).

In vivo implantation of Dex-PBMA

Rat intramuscular and intra-aortic implantations

Dex-PBMA films were first implanted in rat abdominal wall to evaluate the tissue reactions after 7 and 30 days (Fig. 6). As shown on HE staining (Fig. 6A), a tissue surrounded the film at day 7 with a thickness of 40 ± 12 µm that remain stable at day 30 (47 ± 12 µm). The low level of Naphthol AS-D Chloroacetate staining indicated that no neutrophil granulocytes were detected in the tissue (Fig. 6B). The capsule is thus mainly composed by collagen without any evidence of inflammatory reaction or necrosis.

Fig. 6

Histopathologic evaluation for tissue biocompatibility of surrounding tissues after Dex-PBMA intramuscular implantation at 7 and 30 days. (A) Representative photos of hematoxylin-eosin stainings. Double arrows represent cell cap. (B) Granulocytic lineage labeled by naphtol AS-D Chloroacetate esterase staining in purple (arrow). Scale bar 100 μm.

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Fig. 6. Histopathologic evaluation for tissue biocompatibility of surrounding tissues after Dex-PBMA intramuscular implantation at 7 and 30 days. (A) Representative photos of hematoxylin-eosin stainings. Double arrows represent cell cap. (B) Granulocytic lineage labeled by naphtol AS-D Chloroacetate esterase staining in purple (arrow). Scale bar 100 μm.

Dex-PBMA was then evaluated on a coated stent in rat abdominal aorta. Immediately after implantation, no acute thrombosis was observed. After 30 days, all Dex-PBMA or control CoCr stent implanted rats (n=6) survived. At day 30, angiographies confirmed that all arteries were patent. As shown in Fig. 7, no thrombus was found in Dex-PBMA or CoCr stented arteries. Nevertheless, a small intimal hyperplasia was observed both on Dex-PBMA and CoCr stents whereas no significant differences were noticed in the areas of intima (0.54 ± 0.007 mm² vs. 0.50 ± 0.06 mm²) and media (0.40 ± 0.04 mm² vs. 0.43 ± 0.02 mm²) on CoCr stent and Dex-PBMA stent (Fig. 7A). No macrophages were emphasized in the arterial wall (CD 68 staining, Fig. 7B). The neointima was mainly composed of SMCs (α-actin staining). This SMC layer was cover by an endothelial layer (RECA staining) on Dex-PBMA as well on CoCr stents (Fig. 7B). As observed in vitro, no platelets adhered on Dex-PBMA and we noticed the absence of thrombus formation.

Fig. 7

Dex-PBMA coated or bare CoCr stent implantation in rat abdominal aorta after 30 days. (A) Representative Hematoxylin-Eosin stainings. L indicates the lumen, I the intima delimited by the lumen and internal elastic lamina (dashdotted line), M the media delimited by the internal and the external elastic laminas (dashed line) and S a strut of the stent. (B) Representatives α-actin (smooth muscle cells), RECA (endothelial cells) and CD 68 (Macrophages) immuno-stainings. In both cases, a layer of cell proliferated in lumen. It was formed by SMCs and no macrophages were evident. An endothelium was still present. Scale bar 100 μm.

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Fig. 7. Dex-PBMA coated or bare CoCr stent implantation in rat abdominal aorta after 30 days. (A) Representative Hematoxylin-Eosin stainings. L indicates the lumen, I the intima delimited by the lumen and internal elastic lamina (dashdotted line), M the media delimited by the internal and the external elastic laminas (dashed line) and S a strut of the stent. (B) Representatives α-actin (smooth muscle cells), RECA (endothelial cells) and CD 68 (Macrophages) immuno-stainings. In both cases, a layer of cell proliferated in lumen. It was formed by SMCs and no macrophages were evident. An endothelium was still present. Scale bar 100 μm.

Implantation in a rabbit model of intimal hyperplasia

Dex-PBMA or Dex-PBMA _TAC stents were evaluated in a pathological model of neointimal hyperplasia. After 30 days the iliac arteries were patent. Hematoxylin-eosin staining indicated a neointima (Fig. 8A). The measurements emphasized a higher ratio of intima and media areas on Dex-PBMA stent than on CoCr stent (2.39 vs 1.90). Nevertheless, with Dex-PBMA _TAC this ratio decreased (1.60 vs 2.39). On the other hand, TNF-α blood concentration used as a biomarker between before and 30 days after implantation increased in CoCr stent implanted rabbit (+15.6 pg/mL), and remain stable for Dex-PBMA stent implanted rabbit (+0.53 pg/mL) and decreased for Dex-PBMA _TAC stent implanted rabbit (-4.87 pg/mL) (Fig. 8B).

Fig. 8

Dex-PBMA with or without Tacrolimus coated or bare CoCr stent implantation in rabbit iliac artery restenosis model after 30 days (n=1, 3 animals, preliminary study). (A) Representative Hematoxylin-Eosin stainings (Up). L indicates the lumen, I the intima delimited by the lumen and internal elastic lamina (dashdotted line) and M the media delimited by the internal and the external elastic laminas (dashed line). Quantification of intimal hyperplasia (down) on Dex-PBMA (black bar), Dex-PBMA _TAC (grey bar) and CoCr (white) stents. Comparison of areas of the intima and media and the ratio of these last ones. (B) Difference of TNFα arterial blood concentration between before stent implantation and 30 days letter before sacrifice.

×

Dex-PBMA with or without Tacrolimus coated or bare CoCr stent implantation in rabbit iliac artery restenosis model after 30 days (n=1, 3 animals, preliminary study). (A) Representative Hematoxylin-Eosin stainings (Up). L indicates the lumen, I the intima delimited by the lumen and internal elastic lamina (dashdotted line) and M the media delimited by the internal and the external elastic laminas (dashed line). Quantification of intimal hyperplasia (down) on Dex-PBMA (black bar), Dex-PBMA _TAC (grey bar) and CoCr (white) stents. Comparison of areas of the intima and media and the ratio of these last ones. (B) Difference of TNFα arterial blood concentration between before stent implantation and 30 days letter before sacrifice. Fig. 8.

Discussion

Implantable medical devices integration is a challenge. Their biocompatibility relates to the specific interaction between the biomaterial, blood components and tissue vessel. In previous studies, we demonstrated that Dex-PBMA polymer coating improves in vitro the endothelial cell coverage and limits the SMC proliferation involved in the stent restenosis.^22,23 Current complications on stent are not only due to stent restenosis but also to LST.^21,33,34 Both stent geometry and position play a role in stent thrombosis outbreak.^4,35 Stent coating itself is a major factor in its interactions with blood and vascular wall and its quality can be highly variable.³⁵ Indeed, implanted commercial coated stent showed uncoated areas, craters, webbing or cracks.^36,37 Coating integrity and mechanical resistance during implantation and indwell have an impact on long-term stent safety.¹⁴ In the present work, Dex-PBMA stent struts were fully and homogeneously coated. The coating resists to both crimping and stent deployment as well as under the shear stress of physiological flow. This behavior prevents polymer fragmentation and embolization in blood circulation.

Several studies underline the importance of a rapid re-endothelialization of the stented area.¹⁹ In humans, re-endothelialization is due to the proliferation and migration of adjacent endothelial cells (ECs), but also to the contribution of endothelial progenitors cells in the circulation blood that differentiate in ECs.^38,39 For instance, anti-CD34-bio-engineered stents were developed to reduce thrombogenicity.^13,40,41 Because ECFCs that differentiated in vitro from endothelial progenitors and MSCs allows wound healing, we thus investigated their adhesion and growth on Dex-PBMA surface. Measurement of cell adhesion, viability and proliferation revealed no cytotoxic effect.

In previous studies, a lower hydrophobic copolymer of dextran and PBMA (θ = 79.4 ± 0.3 vs θ = 98.5° ± 0.4) improves in vitro the endothelial cell coverage and limits the SMC proliferation involved in stent restenosis.^22,23 This difference could be explained by the control of cerium concentration during the synthesis. Because cerium is able to increase prothrombin formation, coagulation time and induce a proinflammatory cytokine secretion,^42,43 cerium concentration was decreased from 0.2 M to 0.05 M and additional filtration steps were performed. Elementary analysis confirmed the decrease of the residual cerium concentration from 1.73 ± 0.15% to 0.069 ± 0.033%. The hydrophobicity of Dex-PBMA can explain the observed differences in vitro. Numerous works demonstrated the impact of the surface wettability on cell adhesion and proliferation.⁴⁴ Nevertheless, this effect is lowered in vivo, by plasmatic protein deposition, mainly the albumin, which occurred immediately after the device implantation,^45,46 increasing the wettability and favoring the material integration as observed in our in vivo study. Indeed, in both rat healthy aorta and rabbit pathological models, hyperplasia on Dex-PBMA stents were limited and comparable to CoCr Stents. Moreover, immunostaining highlighted a new endothelial cell layer in rats.

As the transcutaneous way of stent implantation may involve a bacterial infection risk and some lethal cases, bacterial adhesion was evaluated.²⁹ However, dextran is known to favor bacterial adhesion and growth,⁴⁷ and functionalized polymethyl methacrylate-based polymers were shown to prevent it.²⁸S. aureus adhesion to Dex-PBMA CoCr disk was thus assessed and was equivalent to the negative control. In contrast, a higher rate of bacterial adhesion was observed on bare CoCr disk, probably due to the electropolishing surface treatment used for clinical stent preparation,⁴⁸ which generates micro-scratches favoring bacterial adhesion on the surface. In our experimental conditions, our Dex-PBMA coating was shown to prevent bacterial adhesion.

Hemocompatibility is identified as a key milestone for a vascular device development. Platelet adhesion and blood coagulation are modulated by material surface.^49-51 As an example, in polylactic acid and magnesium scaffolds or silicone rubbers material used for the biomedical membrane, the thrombosis remains the main complication.^35,52,53 The Platelet adhesion was thus evaluated as reflect of primary hemostasis. We studied either in vitro static and dynamic platelets adhesion as well as thrombin generation on Dex-PBMA surfaces. In static and dynamic conditions few platelets adhered on the surface coated with Dex-PBMA as compared to the bare metal CoCr. We then quantify thrombin generation, which allows us to check the influence of our copolymer on all the coagulation pathways. Dex-PBMA copolymer reduces the formation of thrombin in plasma when coated on CoCr, reflecting an inhibition of the coagulation pathway. Thus, our results showed that Dex-PBMA copolymer is hemocompatible and do not trigger platelets adhesion and activation, as well as blood coagulation.

Moreover in vivo, both in rat and rabbit models, histological results show the absence of acute thrombosis immediately after the implantation or at the time of explantation. Therefore Dex-PBMA exhibits key factors for hemocompatibility.

Since several strategies are currently used to improve the stent efficiency by adding drug in the coating of the whole structure, the abluminal face or by introducing it into the micro-cavities,^11,54 we investigated the possibility to use dextran-PBMA polymer as a drug delivery platform. Tacrolimus, a macrolide molecule loaded in several stents polymer stent coating,⁸ was chosen as a drug model. Under flow, we demonstrated that Dex-PBMA coating could act as a drug eluting platform and release TAC in the medium. A burst release profile (50%) within 3 hours and a steady state over 4 days were observed. These results show a release profile similar to pharmacokinetic profiles of DESs such as the clinical Janus CarbosStentä that releases TAC in the vascular wall tissue within the first hours, and then reaches the steady-state within 6 days.^8,55 The release profile depends on the initial coating as well as the loaded concentration in the stent.⁵⁶ The duration of Tacrolimus release from Dex-PBMA _Tac can be improve using multilayer of Dex-PBMA coating strategy as observed with the Cypher^®21 or the Infinnium^® stents.¹⁵ In our conditions, the low loading level of the drug could explain the results observed in vivo in restenosis rabbit model where no difference was seen between Dex-PBMA _TAC and Dex-PBMA stents. Another study should be devoted in the future to increase the amount of drug and then evaluate its efficiency.

The in vivo implantation of a material and its degradation induce the local repair process, which involves the inflammation and the remodeling process.^57-59 From the in vivo results, neither esterase staining nor TNFα blood cytokine modification was found. No pathologic inflammatory reaction was observed in the intramuscular position or in the aorta. Furthermore, the fibrous cap thickness is important in the case of drug delivery.⁵⁷ A thicker cap decreases drug release. Dex-PBMA shows a moderate thickness of the fibrous cap as compared to other drug delivery platforms.⁶⁰ Therefore Dex-PBMA can therefore be proposed as an interesting local drug delivery platform.

Conclusion

The present study shows that Dex-PBMA exhibited properties suitable for stent coating, resisted to mechanical forces required for crimping coating processes, remained undamaged under flow conditions and could release a drug. Dex-PBMA exhibited moderate platelet adhesion and did not trigger thrombin generation in vitro and allowed endothelial progenitor and MSCs proliferation. In vivo Dex-PBMA i) was efficient to promote endothelial coverage 30 days after stent implantation that might prevent side effects in contact with blood, ii) prevented pathological inflammatory reaction in a rabbit model of intimal hyperplasia. Dex-PBMA represents therefore a new opportunity and an interesting alternative for coating implantable medical devices at the interface with the blood.

Acknowledgments

Cecilia Delattre would like to thank the Groupe de Réflexion sur la Recherche Cardio-vasculaire, the Fédération Française de Cardiologie and the Société Française de Cardiologie for providing a Ph.D. Scholarship. We thank the nursing services of Hôpital St Louis and Bicêtre Hospital for providing umbilical cord blood samples.

Funding sources

This work was supported by the European FP7 program PRESTIGE.

Ethical Statement

The study was approved by the Biological Resources Center in the Cell Therapy Unit from Saint Louis Hospital (Paris, France), authorized by French Ministry of Research under number AC-2008-376 and certified by the French Normalization Agency under number 201/51848.1. The protocol complied with the Declaration of Helsinki. Umbilical cord blood was collected after normal full-term deliveries with the written informed consent of the mother, and used within 24h.

The in vivo procedures and the animal care complied with the “Principles of animal care” formulated by the European Union (Animal Facility Agreement n°75-18-03, 2005) and the protocol was approved by the ethical committee (authorization n°75–214, French Ministry of Agriculture).

Competing interests

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Authors contribution

The conception and design were performed by CD, GPD, CBV, DL and AMP; financial support was obtained by CR; the collection and/or assembly of data was conducted by CD, DV, CR, VO, AL, TA, VG and AMP; data analysis and interpretation were performed by CD, DV, GPD, VO, TA, GC, MJP, CBV and AMP; and the manuscript was prepared by CD, GPD, CBV, DL and AMP; all authors approved and revised critically the manuscript; and given final approval by AMP.

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