HK1126695B - Drug eluting stent with a biodegradable release layer attached with an electro-grafted primer coating - Google Patents

Abstract

The present invention provides a drug eluting stent including a metallic stent framework, an electro-grafted primer coating disposed on the stent framework; and a biodegradable polymer coating hosting a drug disposed on the electro-grafted primer coating and a method of manufacturing said biodegradable drug eluting stent.

Description

Drug eluting stent with biodegradable release layer attached to electro-graft primer layer

Technical Field

The present invention relates to drug eluting stents. More specifically, the invention relates to an adhesive substrate (primer) applied to the surface of a metal stent, which is subsequently coated with a biodegradable polymer capable of containing itself a drug and releasing it in a sustained manner.

Background

The use of coatings for medical devices and drug delivery has been a necessity for many years, notably to increase the capabilities of medical devices and implants. Drug eluting medical device coatings have formed the primary biomedical devices for the treatment of cardiovascular diseases.

Heart disease and heart failure are the two most common health-affecting conditions in the united states and throughout the world. In coronary artery disease, the heart vessels become narrowed. When this occurs, the oxygen supply to the heart muscle is reduced. The main cause of coronary artery disease is the blockage of the artery ("plaque") by fatty deposits. Treatment of coronary artery disease initially used surgery and CABG (coronary artery bypass graft), which is a common and effective method for cardiac surgeons to treat the disease. However, mortality and morbidity are high. In the 60 s, some physicians developed less invasive treatments with medical devices. They can treat the disease by passing through a small incision from the femoral artery: balloon angioplasty (widening of an already narrowed artery using a balloon catheter, which is inflated to open the artery: PTCA ═ percutaneous transluminal coronary angioplasty) is used in patients with coronary artery disease. After balloon angioplasty, within 3 to 6 months, about 40 to 50% of the coronary arteries are affected by restenosis (which narrows again after opening the vessel, usually by balloon angioplasty) due to thrombosis (a blood clot develops in the vessel which can block the vessel and stop the flow of blood) or abnormal tissue growth. Thus, restenosis constitutes a major limitation of the effectiveness of PTCA.

At the end of the 80's, Bare fixed prosthetic valve stents (BMS) were introduced to keep the coronary arteries expanded, indicating some avenues for alleviating this problem and dissecting the arteries while the balloon was inflated in the PTCA procedure. The stent is a mesh tube (a long thin flexible tube that can be embedded in the body; in this case, it can be passed through the heart) fixed to a balloon catheter. But 6 months after stent insertion: the stent is embedded on the side of the growing arterial tissue and the BMS continues to be associated with restenosis at a rate of about 25% in the patient. This tissue consists essentially of smooth muscle cells (SMC's) and, once the stent is placed (in its apposition), the initial injury to the artery stimulates its proliferation. The placement actually destroys the endothelial cell layer (EC's), which has to proliferate and migrate further in order to re-implant (recolonize) onto the stent struts on the SMC's, thereby terminating their proliferation.

The biomedical industry addresses this failure rate in part by designing a new generation of coated stents with the ability to release selective drugs (sirolimus, paclitaxel, ABT578, tacrolimus, everolimus.) in the vessel wall, thereby preventing restenosis. Drug Eluting Stents (DES) have gained increasing attention in the late 90 s as a means of providing more effective reduction of restenosis to a single shape. Theoretically, the drug prevents SMC's from proliferating when active EC's allow early re-transplantation, as later described cells can spontaneously produce Nitric Oxide (NO), a small molecule that serves to terminate SMC's proliferation signals.

Most DES are commercially available as prepared on a polymeric release matrix from which the drug is eluted. The polymer is said to be biostable: the polymer remains permanently on the stent and, therefore, exhibits less effect on both the inflammatory response and the re-transplantation of EC's. The main disadvantage of these DES is that they do not release 100% of the drug they contain. An important consequence of this is that the drug remaining in the coating prevents the re-transplantation process (as most drugs are equivalent to or more effective than "killing" SMC's). This drawback can have fatal consequences for the patient and thus serious consequences for the DES industry. Indeed, although restenosis can be reduced from about 20% with BMS to about 5% with DES, the industry is currently faced with the challenges shown and unresolved by current DES: the subsequent phenomenon of thrombosis, i.e. re-clotting of the arteries one year or more after stent implantation.

It has long been known that implantation of bare metal stents is a source of thrombosis in addition to restenosis, but thrombosis can be readily treated by combining two systemic therapeutic agents, typically aspirin and clopidogrel (Plavix), of two antithrombotic agents). Typically, patients prescribed a stent are administered the two therapeutic agents for 1 to 2 months. Long-term follow-up data indicated that the combination had excellent results on thrombosis. With drug eluting stents, many cases of arterial recondensation have been reported due to clotting (thrombosis) shortly after cessation of administration of both therapeutic agents, which has led cardiologists to advocate the use of the two therapeutic agents for 3, 6, 9, and 12 months or more. Some cases were reported in which myocardial infarction with complete stent thrombosis could occur only a few weeks after 18 months of discontinuation of both therapeutic agents.

Late thrombosis is an acute complication that can be fatal when it occurs if the patient is not in follow-up, or even if he is, while the patient is away from cathlab or a fully equipped medical center. Moreover, both therapeutic agents are very unfortunate constraints, such as that some patients decide themselves to terminate it after an estimated time is long enough, or may forget to take the drug, or may have undergone unexpected clinical intervention, and thus, be in a state where the antithrombotic therapy has to be terminated.

The exact reasons for explaining late thrombosis are still not fully understood. The problem of pathologists evaluating late stage thrombosis revealed that stent re-implantation by EC's was incomplete, leaving metallic or polymeric material in contact with blood for a prolonged period, on which platelet adhesion easily occurs, leading to catastrophic deposition of thrombus. Another explanation arises in the belief that re-implantation of EC's is not entirely a result of incomplete release of the drug from the release layer, where on the surface of the polymer + drug layer it "kills" migrating EC's in their attempted migration and proliferation. The risk of late thrombosis is therefore a serious disadvantage of the existing DES.

Due to the very high mechanical constraints, uncontrolled cracking and delamination of the stent is a common phenomenon (rule) encountered during its manufacturing process (crimping on a balloon), its transfer into the artery (especially calcified lesions) and its expansion (the diameter of the stent increases 3 to 5 times). Cracks and delamination can produce artificial "roughness" ranging from tens of microns to millimeters, and thus, they tend to severely impede proper re-implantation of EC's on stents.

However, "coarseness" alone cannot be interpreted as hindering the re-implantation of EC's alone. One study to evaluate the re-transplantation of EC's was obtained in porcine arteries on day 28 using two overlapping Cyphers in the same arteryOr two TaxusStents, comparing them to their corresponding bare metal counterparts, i.e. using two BxVelocity in the same artery respectivelyOr two ExpressAnd (3) confirming that:

even with DES and BMS (because of the overlap between the two stents), the "roughness" of the surface of the re-implant is rather high, and re-implantation with BMS is always better compared to DES.

Irrespective of DES, CypherOr TaxusI.e. no matter what drug is being released, the re-transplantation with the corresponding BMS is always better.

This result strongly suggests that after the coating is deployed, in addition to the "roughness" of the coating and the stent surface, re-implantation is always better than in the absence of the drug. This is related to the fact that:

all DES present have a biostable layer. The release of the drug is obtained by pure diffusion and therefore will never be complete: there is always some drug remaining in the coating to be re-implanted for the duration of the period;

all drugs used in the existing DES (sirolimus, paclitaxel) have threshold toxic concentrations that have comparable or even lower resistance to EC's than SMC's, i.e. they can "kill" EC's very equally or even more than SMC's:

this points to a serious disadvantage of existing DES, which locally retain drugs that are toxic to EC's for extended periods of time.

Last but not least, these drugs may have an effect on the remodeling of the arteries. It is noteworthy that the so-called "stent malaposition" shows that some stent struts are not completely in contact with the arterial wall. It is believed that most stents are due to the effect of the drug, particularly in the case of sirolimus, which causes so-called arterial "positive remodeling", i.e. its gradual over-expansion (over-dilation): the stent initially makes good contact with the arterial wall but eventually "floats" within the artery, increasing in diameter under the effect of the drug. In this case, some stent struts remain un-reimplanted by EC's (since they are very far from the arterial wall) and can be a source of thrombus formation that is blocked from direct contact of the polymer substance with blood. These thromboses may not appear as long as the patient is under both anti-platelet therapeutics, but will begin to form immediately after both therapeutics are discontinued (late thrombosis). This in turn points to a serious disadvantage of the existing drug eluting stents, due to the long retention of the drug on the stent surface.

Disclosure of Invention

It is an object of the present invention to provide a stent with short term DES-like properties to prevent restenosis and a long term BMS-like history to avoid thrombosis and allow EC's to proliferate and migrate early before remodeling. As detailed earlier, late thrombosis is believed to involve:

incomplete drug release;

poor coating integrity due to lack of adhesion of the coating on the stent surface, leading to cracks and delamination, a potential source of "roughness" that prevents re-implantation by EC's;

poor healing-promoting (long-term) behavior of coatings for EC re-implantation in case of poor stent placement (ISA) due to drugs.

In the DES according to the invention, the drug, if any, is released by the biodegradable polymer, which will disappear after several weeks, thereby releasing 100% of the drug. The preparation of coatings using these biodegradable polymers must be carried out by using an adherent underlayer in order to significantly promote good mechanical integrity for proper stent placement.

Accordingly, a drug eluting stent is suggested, comprising:

a stent frame;

an electro-graft coating applied to the stent framework, and

a biodegradable polymer coating containing a drug coated on the electro-implant coating.

The DES may further comprise a biodegradable topcoat.

The electro-graft coating is used as an effective substrate coating that promotes adhesion between the surface of the metal stent and the subsequent polymer coating. The electro-graft coating may be applied to the stent, dried, and then the drug polymer applied. The subsequent polymeric coating may comprise one or more therapeutic compounds that provide drug properties to the drug eluting stent. The substrate electrotransport coating acts as a bridge between the substrate and the organic polymer coating and has good adhesion to metals and drug polymers.

The electro-grafting technique allows covalent bonding on the surface, giving the layer tens to hundreds of nanometers, and the deposition of nanometric controls (nanometric controls), as well as known vascular materials such as p-BuMA. Moreover, the obtained electro-grafted layer is homogeneous and conformal to the stent surface (conformal). The electrographic coating (i) disappears itself, i.e., is itself biodegradable; or (ii) exhibit good properties with respect to cell migration and proliferation, in particular absolute homogeneity and absence of cracks and delaminations. When the biodegradable release matrix disappears, the underlayer will contact the EC's or SMC's (undergoing re-transplantation) or blood (incomplete re-transplantation, ISA, etc.), or both. It is therefore particularly important that the underlayer itself be as homogeneous as possible, in particular that it be crack-free, which would prevent complete re-implantation with EC's.

Drawings

Fig. 1(a) shows the cumulative release of sirolimus over time (days) from PLGA bilayer coatings in vitro.

Fig. 1(B) shows the cumulative release of sirolimus over time (days) from the poly (lactide) bilayer coating in vitro.

Fig. 2 shows the fractional release of sirolimus from PLGA or PLA over time (days) in vivo.

Detailed Description

A first object of the invention is a Drug Eluting Stent (DES) comprising:

a stent frame;

an electro-graft coating applied to the stent framework, and

a biodegradable polymer coating containing a drug coated on the electro-implant coating.

Support frame

Wherein the stent framework advantageously comprises a metal matrix. In particular, the scaffold frame comprises a substance selected from the group consisting of: stainless steel, nickel, tantalum, cobalt-chromium MP35N or MP20N alloy, platinum, titanium, a suitable biocompatible alloy, a suitable biocompatible substance, and combinations thereof.

Electro-implant coating

The electro-graft layer serves as a substrate for the overlying biodegradable layer (during the preparation, crimping and scaffolding process). The coating of the electrotransport substrate is a uniform layer. The layer preferably has a thickness of 10nm to 1.0 micron, in particular a thickness of 10nm to 0.5 micron, more in particular 100nm to 300 nm. Such a thickness, which is smaller than the minimum radius of curvature achievable at any point on the stent surface, ensures that the coating does not crack. The electro-grafted layer can prevent cracking and delamination of the biodegradable polymer layer and shows equal or better re-grafting than the stainless steel BMS. Moreover, the use of an electro-grafted layer having a thickness of at least about tens or one hundred nanometers ensures a good enhancement of the adhesion of the overlying biodegradable layer due to the interdigitation between the two polymeric layers. In this sense, the properties of the electro-implant polymer are selected based on the properties of the release matrix polymer, which itself is selected based on the loading and kinetics of the desired drug release: the electro-implant polymer and the release matrix polymer must be partially miscible to make a good interface. This is the case: for example when the two polymers have similar solubility or Hildebrand parameters, or when the solvent of one of the polymers has at least good swelling properties (swellan) for the other polymer. In addition to being bound by these, the properties of the electro-grafted polymer are preferably from the polymer list of known biocompatibility. Finally, not all polymers can be obtained by electro-grafting, but most polymers obtained by the propagation chain reaction are acceptable, such as vinylates (vinyls), epoxides, cyclic monomers via ring-opening polymerization. Thus, polybutylmethacrylate (p-BuMA), Polymethylmethacrylate (PMMA) or PolyEpsilonCaprolene (p-ECL) are polymers of interest, obtainable by electro-grafting, which interact with hydrophobic release matrices. Polyhydroxyethylmethacrylate (p-HEMA) is a polymer of interest, obtainable by electro-grafting, which interacts with a hydrophilic release matrix.

Other organic films, obtainable with electro-grafting, but which do not have "true" polymeric properties, can be very effective substrate layers for the release matrix: this is the case for the "poly" nitrophenyl films obtained, thanks to the electro-grafting of phenyl diazonium salts, in particular 4 aminophenyl diazo tetrafluoroborate, on the surface of said scaffolds before the matrix is sprayed for release. The phenyldiazonium salt is preferably of the formula Y-ArN₂ ⁺X-, wherein Ar represents an aryl group, advantageously a phenyl group, X represents an anion, advantageously selected from: halogen ion, sulfate ion, phosphate ion, perchlorate ion, boron tetrafluoride ion, hexafluorophosphate ion and carboxylate ion, Y is a functional group, and is advantageousIs selected from: nitro, hydroxyl, thiol, amino, carboxyl, carbonyl, ester, amino, cyano, alkyl or functionalized alkyl, phenyl or functionalized phenyl.

The electro-grafting layer, in particular the p-BuMA layer, may further have a passivating behavior and block the release of heavy metal ions (in the blood flow or in the arterial wall) from the stainless steel surface. The heavy metal ions are believed to be responsible for the initiation of inflammation caused by the introduction of the metal scaffold into the blood, which is an electrolyte mediator and thus causes partial oxidation of any metal until the Nernst equilibrium is reached. In particular, the thickness of the arterial wall of the electro-grafted and biodegradable (drug-free) layers in the study, observed from a longitudinal section, was always less than that of the bare metal scaffold layer, demonstrating less granuloma, i.e. less inflammation: this result was confirmed by the results observed with the 28-day rabbit study, in which less inflammation was detected with a stent coated with only the electroporated p-BuMA layer, as compared to BMS (see examples 11 and 12).

In one embodiment of the invention, the electro-graft layer itself is biodegradable, so that it disappears from the surface of the stent after the biodegradable release layer has also disappeared.

The electro-graft layer has a non-thrombogenic (or anti-thrombogenic) effect and a pro-healing effect (promoting proliferation and adhesion of active EC's once the biodegradable release layer has disappeared). The mechanism of hydrolysis of the biodegradable polymer will not continue until the EC's begin to proliferate in the drug-containing biodegradable layer, i.e., before it completely disappears, and the EC's will soon come into contact with the electro-grafted layer. If the electro-graft layer is itself biodegradable, the healing promoting effect is expected to be of a stainless steel surface. Greater healing promoting effects are obtained with biostable electro-grafted layers, which ensure proper re-implantation with EC's over a longer period of time.

A 60 day pig trial was performed using a composition layer consisting of a bottom layer (150nm) of electro-grafted p-BuMA (polybutylmethacrylate) and a biodegradable release layer (5 μm) of PLGA (polylactide-co-glycolide) coated thereon, and described in example 13. The results of this study will first show that the biodegradable release layer disappears after the first 4 weeks, thereby releasing 100% of the drug. It was also demonstrated that at 8 weeks, the stent coated with the electro-graft layer and the biodegradable layer was completely re-implanted by endothelial cells; since the biodegradable layer is known to disappear after 4 weeks, this means that good replanting is the result of the electro-graft layer alone interacting with the artery and blood stream.

The combined properties of the electro-graft layer and the biodegradable scaffold were statistically higher than BMS, even in (difficult) locations of two layers of the composition (electro-graft layer + biodegradable reservoir) where there was no drug inside the biodegradable release layer. The DES according to the invention, shortly after implantation of the stent, will be able to terminate both antiplatelet therapies due to better re-transplantation by EC's.

Thrombosis is a phenomenon initiated by the adhesion of specific proteins on surfaces to which the antithrombotic behaviour relates to minimizing or even eliminating the tendency of proteins to adsorb. Several types of polymers having this anti-fouling effect are known in the art, such as heparin, CMDBS, PC (phosphorylcholine) -based polymers, etc., more typically polymers carrying zwitterionic groups, polyethylene oxide (PEO) or polyethylene glycol (PEG), and more typically polymers having an arbitrary highly hydrophobic surface. These polymers have the common feature that they carry very little, if any, active functionality that tends to facilitate protein binding to their surface.

In short, the electro-graft layer may additionally be composed of these anti-fouling substances, provided that they are also compatible with the aforementioned criteria, which have a good interface with the release matrix polymer in order to have an acceptable antithrombotic behaviour. This condition is not in contradiction to the properties, the electro-graft layer acting as a substrate layer, must be conformed such that it adheres an increasingly thick biodegradable layer to the metallic surface of the stent, as we see above, since the adhesion to the release matrix polymer is mainly due to the interdigitation with the electro-graft polymer. It will be noted that the PC polymers studied by the biocompatible Plc are ethylene polymers and can therefore be obtained by electro-transplantation (p-MPC/BUMA, p-MPC/DMA/TMSPMA, see below).

Among the polymers which can be used as electro-implant coatings, mention may be made in particular of vinyl polymers such as: polymers of acrylonitrile, methacrylonitrile, methyl methacrylate, ethyl methacrylate, propyl methacrylate, butyl methacrylate, hydroxyethyl methacrylate, hydroxypropyl methacrylate, cyanoacrylate, acrylic acid, methacrylic acid, styrene and its derivatives, N-vinylpyrrolidone, vinyl halide, and polyacrylamide; isoprene, ethylene, propylene, ethylene oxide, molecules containing cleavable rings such as lactones, in particular epsilon caprolactone, lactide, glycolic acid, ethylene glycol, polyamides, polyurethanes, poly (orthoesters) and polymers of polyaspartic acid.

The organic film obtained by electro-grafting may be a vinyl polymer or copolymer, in particular polybuma (polybutylmethacrylate), polyhema (polyhydroxyethylmethacrylate), poly MPC/BUMA (poly 2-methacryloyloxyethylphosphorylcholine/butyl methacrylate) and poly MPC/DMA/TMSPMA (polymethacrylyloxyethylphosphorylcholine/lauryl methacrylate/trimethylsilyl propyl methacrylate). In one embodiment, the organic film is a biodegradable polymer, in particular polycaprolactone, Polylactide (PLA) or polyglycolide (plga).

Adhesion between an electro-implantable coating and a biodegradable layer (comprising a drug layer or an overcoat layer)

The above biodegradable layer can be adhered to the electro-grafting layer, and pass through;

formation of chemical bonds with the electro-grafted polymer (see, for example, patent application WO04/005410, incorporated herein by reference);

inserting the precursor of said biodegradable layer into the electro-grafted polymer chemical so as to cause its formation inside the electro-grafted polymer film, which will then act as a fixing layer for said biodegradable layer (see for example patent applications WO04/074537 and WO04/075248, incorporated herein by reference);

interpenetration of the pre-formed biodegradable polymer within the electro-grafted layer by cross-bonding. Interdigitation involves that the polymeric chains of the biodegradable polymer may "creep" or "reptate" within the electro-implanted layer and form at least one "loop" within the electro-implanted layer. For polymers, one "loop" is the typical size of a chain when randomly configured; it can be evaluated by the so-called radius of gyration of the polymer. Even though it has some relationship with the specific polymer, its molecular structure, etc., the radius of gyration of the polymer is mostly less than 100nm, indicating that the electrical graft layer must be thicker than the threshold value of the ring shape of the polymer which can contain at least one constituent upper layer, for improved adhesion.

Cross-bonding is a method to obtain excellent adhesion of the biodegradable layer on the electro-grafted layer, as long as the subsequent conditions (visualized the later) are met:

thicker than about 100 nm;

have the same wettability (i.e. hydrophobicity/hydrophilicity) as the above biodegradable polymer, to enable "mixing" between the two;

have a glass transition temperature less than that of the above biodegradable polymer in order to obtain thermal cross-bonding at a temperature sufficiently low to maintain the stability of the drug; or

Is swollen by at least a solvent capable of dissolving the above biodegradable polymer or a solvent of a dispersion containing the above biodegradable polymer or a component thereof, so that the interdigitation can be efficiently formed at room temperature by the addition of droplets only from these liquids on the surface of the electro-grafted layer: the liquid, which swells the electro-graft layer, causes components of the solution or dispersion to intercalate into the electro-graft layer, which is then evaporated to form the interleaved composite material.

This subsequent condition is sufficient to form a good and stable joint between the biodegradable release layer and the electro-grafted layer.

Compared to chemical bonding or layering, interdigitation is a preferred embodiment of the invention to establish a junction between the release matrix and the electro-grafted layer: since the release matrix chosen herein is biodegradable, what remains is a polymer of known structure (electro-grafting), i.e. without unreacted chemical groups or hydrolysable bonds that can promote the reaction of the remaining electro-grafted membrane, a reaction that is liable to cause inflammation and/or thrombosis.

Staggered bonding requires that one can coat a solution containing a biodegradable polymer layer and optionally a drug, optionally selectively with the desired wettability, on the stent coated with the electro-graft layer. For example, PLGA is readily soluble in dichloroethane, dichloromethane or chloroform and is useful for most hydrophobic drugs such as sirolimus, paclitaxel or ABT-578. In this case, electro-implanted p-BuMA is a suitable choice with the desired cross-linked surface because it is readily swollen by (and even soluble in) chloroform or methylene chloride.

From a production point of view, the coating can be done by dipping or by spraying. Less impregnation is used because it does not enable one to obtain layers thicker than about 2 to 3 μm per shot (shot): for greater thicknesses, one must completely dry the first layer before re-impregnation to avoid re-dissolving the already applied layer. This binding makes impregnation very inconvenient for layers exceeding 2 μm. In this respect, spraying is easy to carry out (see example 14). The nozzle spraying the above solution faces the support and rotates it so as to be present on all the external surfaces of the spray. In order to obtain a proper cross-joint surface under the above conditions, one should advantageously introduce so-called "wet spray" or "low pressure" conditions: the solution to be sprayed has a low viscosity (typically < 1cP, pure chloroform with a viscosity of 0.58cP), a short distance from the rotating support, and a nozzle in which the pressure of an inert carrier gas (nitrogen, argon, compressed air) is typically less than 1 bar. These conditions result in the liquid being sprayed into fine droplets, which move in the atmosphere of the spray chamber, causing them to hit the electro-graft scaffold: since the electro-grafted polymer layer and the spray solution have the same wettability, the droplets show a very low contact angle (good wetting), and therefore the droplets pool on the surface and thus form a film (filmogenic) at an early stage. In addition to making a good joint between the biodegradable layer and the electro-graft layer, a "low-pressure" spray device enables the preparation of coated stents with a very small net shape between the struts.

The relative movement of the nozzle and the stent can precipitate a uniform and relatively thin (< 1 μm) layer in a single hit, which is still filled with solvent. Rotation and air renewal can evaporate the solvent, making it easier due to the thin layer, leaving a polymer layer (+ drug) on the surface. Then, a second layer is sprayed on the first layer, and so on, in order to achieve the desired thickness (and thus loading). Since it applies several sprays in order to achieve the desired thickness, the "low-pressure" spray apparatus can be used in batches, with several holders rotating in parallel with one spray nozzle for each holder, thereby enabling the other holders to be evaporatively dried while the other holder is spraying. This can keep the throughput of the system high enough even if the low pressure spraying process itself is highly continuous.

These low pressure spray systems are listed in example 14 and can process 20 rotating carriers per batch with a single nozzle moving on the carrier as controlled by the X-Y scanning system process. One characteristic of this system is that when the X-Y system is outside the box, the rotating support is in the box (able to extract the solvent, and safe for the operator): the movement of the nozzle is guided through the roof of the box by a magnet, holding the "closed shell" structure of the box so that the gloved sample rack is tucked therein from the side door by means of a movable sample-carrier and inserted and connected to a spindle in the box when the door is opened from the inside.

Biodegradable layer comprising a drug

The biodegradable release layer will advantageously have a thickness of 1 to 200 μm, more advantageously about 1 to 10 μm (depending on the load), in order to obtain release of the drug over a defined period of time.

The drug polymer coating may include one or more drugs. Each drug may include a bioactive agent. The bioactive agent may be a pharmacologically active drug or a bioactive compound. The drug polymer coating may degrade during preparation, packaging, sterilization or storage of the drug polymer eluting stent. During sterilization, for example, there may be oxidation of the drug or polymer, resulting in hydrolytic damage, cleavage of polymeric bonds, polymer and/or drug destruction, or actual cracking or spalling of the drug polymer coating. Temperature deviations in the prepared or prepared stent can cause delamination of all or a portion of the drug polymer coating. The present invention solves this problem by using an electrotransport substrate coating between the polymer-drug coating and the metal stent, thereby reducing or preventing drug-polymer delamination.

The drug may be encapsulated in the drug polymer coating using albumin, liposomes, ferritin or other biodegradable proteins and phospholipids, using microbead, microparticle or nanocapsulation techniques, prior to application to the substrate coated stent.

Wherein the bioactive agent may include an anti-neoplastic agent such as triethylenethiophosphoramide, an antiproliferative agent, an antisense agent, an antiplatelet agent, an antithrombotic agent, an anticoagulant, an antibiotic, an anti-inflammatory agent, a gene therapy agent, an organic drug, a pharmaceutical compound, a recombinant DNA product, a recombinant RNA product, collagen, a collagen derivative, a protein analog, a saccharide derivative, or a combination thereof.

The bioactive agent can be any therapeutic substance that provides therapeutic properties for the prevention and treatment of a disease or disorder. The antineoplastic agent can prevent, kill or block the growth and spread of cancer cells in the vicinity of the stent. Antiproliferative agents may prevent or stop cell growth. Antisense agents can be used at the genetic level to interrupt the process of producing disease-causing proteins. The antiplatelet agent can act on platelets to inhibit the blood coagulation function. Antithrombotic agents can actively delay clot formation. Anticoagulants can delay or prevent blood coagulation, and compounds such as heparin and coumarin can be used for anticoagulant therapy. Antibiotics kill or inhibit the growth of microorganisms and are useful in combating diseases and infections. Anti-inflammatory agents may be used to counteract or reduce inflammation in the vicinity of the stent. Gene therapy agents can alter the expression of human genes to treat, cure, or ultimately prevent disease. The organic drug may be any small molecule therapeutic substance. The pharmaceutical compound may be any compound that provides a therapeutic effect. The recombinant DNA product or recombinant RNA product may include modified DNA or RNA genetic material. Pharmaceutically useful bioactive agents may also include collagen and other proteins, carbohydrates and derivatives thereof. For example, the bioactive agent may be selected for inhibiting vascular restenosis, a condition that corresponds to narrowing or contracting the diameter of the body lumen in which the stent is placed. The bioactive agent generally controls cell proliferation. Controlling cell proliferation can include increasing or inhibiting growth of a target cell or cell type.

The bioactive agent may be an agent against one or more of the following conditions: including coronary restenosis, cardiovascular restenosis, angiographic restenosis, arteriosclerosis, hyperplasia and other diseases and conditions. For example, the bioactive agent may be selected for inhibiting or preventing vascular restenosis, a condition that corresponds to narrowing or contracting the diameter of the body lumen in which the stent is placed. The bioactive agent generally controls cell proliferation. Controlling cell proliferation can include increasing or inhibiting growth of a target cell or cell type.

The bioactive agent may include podophyllotoxin, etoposide, camptothecin analogs, mitoxantrone, sirolimus and derivatives or analogs thereof. Podophyllotoxin is an organic highly toxic drug that has anticancer properties and inhibits DNA synthesis. Etoposide is a semisynthetic antineoplastic agent derived from podophyllotoxin, and can be used for treating monocytic leukemia, lymphoma, small cell lung cancer and scrotal cancer. Camptothecin is an anticancer drug that can act as a topoisomerase inhibitor. Camptothecin analogs structurally related to camptothecin, such as aminocamptothecin, are useful as anticancer agents. Mitoxantrone is also an important anticancer drug, commonly used to treat leukemia, lymphoma, and breast cancer. Sirolimus is a drug that interferes with the normal cell growth cycle and can be used to reduce restenosis. The bioactive agents may also include analogs and derivatives of these agents. Antioxidants may be beneficial in themselves due to their antimistenic properties and therapeutic effects.

The drug polymer coating may soften, dissolve or erode from the stent to elute the at least one bioactive agent. This elution mechanism may be referred to as surface erosion, wherein the outer surface of the drug polymer coating dissolves, degrades, or is absorbed by the body; or skeleton erosion (bulk oxidation) in which a substantial portion of the drug polymer coating biodegrades to release the bioactive agent. The eroded portion of the drug polymer coating may be absorbed, metabolized, or otherwise removed by the body.

The drug polymer coating may also include a polymer matrix. For example, the polymer matrix may include caprolactone-based polymers or copolymers, or various cyclic polymers. The polymer matrix may comprise various synthetic and non-synthetic or naturally occurring macromolecules and derivatives thereof. The polymer is advantageously selected from one or more biodegradable polymers, such as polymers, copolymers and block polymers, in various combinations. These biodegradable (also referred to as bio-resorbable or bioabsorbable) polymers include polyglycols, polylactides, polycaprolactones, polyglycerol sebacates, polycarbonates such as tyrosine derived, biopolyesters such as poly (beta-hydroxyalkanoates) (PHAs) and derived compounds, polyethylene oxides, polybutylene terephthalate (polybutylene terephthalate), polydioxanone, hybrids, compositions, collagen matrices with growth regulators, proteoglycans, glycosaminoglycans (glycosaminoglycans), vacuum formed SIS (small intestinal submucosa), fibers, chitin and dextran Polyglycolic acid (PGA) polymer, poly (e-caprolactone) (PCL), polyacrylate, polymethacrylate, or other copolymers. The drug may be dispersed throughout the polymer matrix. The drug or bioactive agent may diffuse out of the polymer matrix to elute the bioactive agent. The drug may diffuse out of the polymer matrix and into the biomaterial surrounding the stent. The bioactive agent can be separated from the drug polymer and diffuse out of the polymer matrix into the surrounding biological material. In a further embodiment, the pharmaceutical coating composition may be prepared using the drug 42-Epi- (tetrazolyl) -sirolimus, as described in U.S. Pat. No.6,329,386 to Abbott Laboratories, Abbott Park, I11, and from dispersed within a prepared (coated) coating by coating with choline phosphate of the Biocompatibles International P.L.C., as described in U.S. Pat. No.5,648,442.

The polymer matrix may be selected to provide a desired elution rate of the bioactive agent. The drug may be synthesized so that a particular bioactive agent also has two different elution rates. For example, a bioactive agent with two different elution rates will allow for rapid delivery of the pharmacologically active drug within twenty-four hours of surgery, while slow, stable delivery of the drug within, for example, two to six months thereafter. The electrografting substrate coating can be selected to stably immobilize a polymer matrix comprising a rapidly releasing (deployed) bioactive agent and a slowly eluting drug on the stent framework.

Outer biodegradable layer

The DES may further comprise an outer coating (topcoat layer) prepared from the same composition as the biodegradable coating release layer. In particular, the outer biodegradable layer may comprise a biodegradable polymer such as Polylactide (PLA), polyglycolic acid (PGA) polymer, poly (e-caprolactone) (PCL), polyacrylate, polymethacrylate, or other copolymers.

Preparation method

Electro-implantation of polymers is a technique based on the formation of a polymer layer on an in situ surface, i.e. from a precursor water bath instead of a pre-prepared polymer. The surface of the coating is electronically polarized and functions as a polymerization initiator, which causes surface polymerization by propagating chain reactions (see FR 2821575; incorporated herein by reference).

The present invention uses a mode of operation in which it is possible to easily carry out the actual polymer electrotransport starting from a precursor solution which is easy to prepare and control, in particular:

(i) a test design that applies an electrode potential that drives the transplantation reaction;

(ii) an electrolytic medium is used which is at least a good swelling agent for the polymer formed, or a good solvent for said polymer.

A biocompatible adhesive film (e.g. polybutylmethacrylate (p-BuMA)) can be obtained on a voltammetric (voltammetric) scanning stent (stainless steel, chromium cobalt alloy) under the following conditions: in a solution (solvent. RTM. DMF) comprising a diazonium salt (especially an aryl diazonium salt such as 4-nitrodiazobenzenetetrafluoroborate) and a monomer (p-BuMA, 3.5mol/l), the diazonium salt is present at a concentration of 5.10^-4To 10^-1Mole/l (especially 10)^-2Mole/l), possible rangesThe sweep rate was 100mV/s, with a range of-0.2V/ECS to 3.0V/ECS.

The electrolytic solution may include a solvent that is intended to solubilize the chain polymerizable monomer primarily by the practitioner (i.e., without intervening electropolymerization reactions). However, the monomers may act as a solvent, such that the presence of these liquids is not necessarily required. When they are used, these solvents are preferably selected from the group consisting of dimethylformamide, dimethyl sulfoxide, ethyl acetate, acetonitrile, tetrahydrofuran, propylene carbonate and other solvents commonly used in electrochemistry, dichloroethane and common chlorinated solvents. The solvent may also be selected from water and alcohols. It is not necessary to subject the reaction medium to distillation prior to dissolution in order to remove the water they contain, nor to strictly control the water content of the atmosphere above the reaction medium. Thus, the method can be easily implemented on a large scale.

The electrolyte may also comprise at least one supporting electrolyte in order to ensure and/or improve the current passing in the electrolyte. When they are used, the supporting electrolyte is preferably selected from quaternary ammonium salts such as perchlorate, tosylate, tetrafluoroborate, hexafluorophosphate, quaternary ammonium halides, sodium nitrate and sodium chloride. The electrolyte may further contain an agent (surfactant) for improving the uniformity of the film, such as glycerin.

The film, if present, has little crosslinking and its adhesion to the surface is a result of bonds with the underlying metal. For this reason we shall use the term electro-grafting of polymers, hereafter even if it now refers to grafting obtained by electro-reduction of a solution comprising monomers that can undergo a propagation chain reaction and a diazonium salt, the latter preferably in low concentration. Such a method enables electro-grafting on all conductive substrates such as organic thin film scaffolds, in particular polymeric thin films having a thickness of several tens to several hundreds of nanometers.

The electro-graft solution coated on the stent framework is dry. Excess liquid may be blown off before drying the film. Drying may be carried out at room temperature or elevated temperature in a dry nitrogen or other suitable environment, including a vacuum environment, in order to eliminate or remove any volatile components from the polymerization solution. The coated stent can be dried at moderate elevated temperature for 60 minutes at 40 ℃ and under vacuum (-10 mbar) to remove any solvent trapped inside the substrate coating. The thickness of the coating of the electro-implant substrate may be in the range of 10nm to 1.0 micron in order to adequately coat the stent framework and provide a satisfactory underlayer for subsequent application of the drug polymer. Additional application and drying steps may be included to achieve the desired thickness of the substrate coating.

For the coating of the electro-grafted substrate, the wet process may be carried out by spraying or by dipping. The drug polymer may be mixed into a suitable solvent and applied to the substrate using an application method such as dipping, spraying, painting or brushing. The drug polymer adheres well to the coating of the electro-grafted substrate during the coating operation. The drug polymer coating may be applied immediately after the coating of the electro-grafted substrate is applied. Alternatively, the drug-polymer coating may be applied to the stent with the coating of the electro-grafted substrate.

The drug polymer may be mixed with a suitable solvent to form a polymeric solution. The drug polymer may include a polymer matrix and one or more therapeutic compounds. To form the drug polymer coating, monomers such as vinyl acetate derivatives can be mixed with other monomers in a solvent such as isopropanol to form a polymeric solution. The mixture may be reacted to form a polymer, and one or more bioactive agents may be mixed with the polymerized mixture to form a drug polymer having a predetermined elution ratio. A suitable bioactive agent or solution comprising a bioactive agent may be mixed with the polymerization solution. Alternatively, a polymer such as a copolyester or block copolymer may be dissolved in a suitable solvent, and one or more bioactive agents may be added to the mixture. The mixture may be mixed with an adhesion promoter in the polymerization solution. One or more adhesion promoters are selected and added to the mixture.

The polymerization solution may be applied to the stent framework with the coating of the electro-grafted substrate. The polymeric solution may be applied to the stent using any suitable method for applying a polymeric solution.

Excess liquid may be blown off and the polymerization solution dried. Drying may be carried out at room temperature or elevated temperatures (40 ℃) and under dry nitrogen or a suitable environment to remove any volatile components of the polymerization solution. Another dipping and drying step may be used to thicken the coating. The thickness of the drug polymer coating may be between 1.0 micron and 200 microns or more in order to provide sufficient and satisfactory pharmacological benefit with the biologically active agent.

The treatment of the drug polymer coating may include air drying or cryogenic heating in air, nitrogen or other controlled environment. The drug polymer coating may be treated by heating the drug polymer coating to a predetermined temperature.

More specifically, illustrative examples of the invention are provided herein. The following examples illustrate:

(1) electro-grafting solution formulation

(2) Method for electro-grafting on stainless steel stent

(3) Electro-grafting method on cobalt-chromium stent

(4) Erosion Barrier Effect of electro-transplanted p-BuMA

(5) Erosion Barrier Properties of Dip coated coupons (dip coated coupons) for electro-grafting p-BuMA and PLA

(6) Spraying method for depositing a reservoir layer

(7) Enhanced adhesion with electrotransport layer

(8) Examples of in vitro drug release kinetics

(9) Cytotoxicity Studies of electrotransport coatings

(10) Hemolysis study of the electro-grafted coating

(11) Local tolerance study of electro-grafted scaffolds after local implantation

(12) Re-implantation properties of the BuMA coated stents compared to BMS at 14 and 28 days in the rabbit model

(13) Local tolerability behind fully coated stents in swine

(14) Low pressure spray system for preparing DES with good interface with electro-grafted layers

Example 1: electro-grafting solution formulation

One embodiment of the invention is a formulation of an electrografting solution based on the vinylic monomer n-butyl-methacrylic acid (BuMA) dissolved in DMF solvent. NaNO₃Is used as an electrolyte support.

Table 1: electro-grafting solution formulation

Example 2: method for electro-grafting on stainless steel stent

The p-BuMA coated 18mm stainless steel crown stent (ClearStream Technologies) was electro-grafted with the chemical solution described in example 1 with the following parameters, rinsed and dried at 40 ℃ for 60 minutes under a vacuum of 10 mbar. The coating thickness obtained with this method was about 150 nm.

Electrical transplantation parameters:

the method comprises the following steps: cyclic voltammetry with a break potential of-3.2V/CE and argon (2 Lmin) injection^-1)。

The scanning times are as follows: scanning 50 times

Scanning rate: 50m V/s.

Example 3: electro-grafting method on cobalt-chromium stent

Using the chemical solutions described in example 1, a 18mm cobalt-chromium coronary stent (Natec-median) coated with p-BuMA was electro-grafted with the following parameters, rinsed and dried at 40 ℃ for 60 minutes under a vacuum of 10 mbar. Before electro-transplantation, with NH₄The stent surface was treated with the F40% solution for 1 minute. The coating thickness obtained with this method was about 150 nm.

Electrical transplantation parameters:

the method comprises the following steps: cyclic voltammetry with a break potential of-3.5V/CE and argon (2 Lmin) injection^-1)。

The scanning times are as follows: scanning 50 times

Scanning rate: 50 mV/s.

Example 4: erosion Barrier Properties of electro-grafted p-BuMA

The corrosion resistance of the electro-grafted p-BuMA was evaluated on coated stainless steel coupons synthesized according to the experimental design described in example 2.

For this purpose, the p-BuMA coated samples (test) and the uncoated samples (control) were electroporated to a value of 1cm²9g/l of surface area/volume ratio/ml were immersed in NaCI solution. The samples were kept at 37 ℃ with gentle agitation and the time course of cobalt, nickel and molybdenum ion evolution was evaluated by taking samples of the release medium as usual. The ions were quantified using an inductively coupled plasma-mass spectrometer (ICP-MS).

Table 2: ion release

The release of ions from the metal surface is strongly reduced by electro-implantation of a p-BuMA coating, e.g., released nickel (to date)So far, it is the most toxic element) from 28ng/cm on a stainless steel specimen²Reduced to 3ng/cm on an electro-grafted p-BuMA tablet²。

Example 5: erosion Barrier Properties of electro-grafted p-BuMA and PLA Dip smears

Stainless steel electro-grafted p-BuMA tablets were dip-coated in polylactide (p-PLA) solution (5% w/v in chloroform) with or without 20% (w/w) of the model drug pentoxifylline. After immersion, the coating was stabilized at room temperature for 24 hours and dried in an oven at 40 ℃ for 48 hours. Ion release was performed according to the experimental design described in example 3. Comparison of the release of Cr (A), Ni (B) and Mo (C) ions in 150 days from 316 stainless steel coupons, electrografted p-BuMA + p-PLA dip coupons and electrografted p-BuMA + p-PLA dip coupons containing pentoxifylline in a solution of NaCl (9g/l) is given in Table 3 below:

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TABLE 3

The same reduced ion release was observed on stainless steel coupons coated with a double layer coating, with the content of Ni ions being significantly from 30ng/cm²Reduced to about 8ng/cm²And, for the electro-grafted p-BuMA/p-PLA coating and the electro-grafted p-BuMA/PLA/pentoxifylline coating, respectively, the Cr ion content is from 18ng/cm²Reduced to 6ng/cm²And 4ng/cm²。

Example 6: spray coating method

A spray coating process for applying a depot polymer coating to an electro-grafted metal stent shows another embodiment of the invention. After drying, an 18mm electro-graft stent was spray coated with a biodegradable polyester containing sirolimus (polylactide-co-glycolide 50/50, PLGA).

The copolymer (0.25% w/v) was dissolved in chloroform. Sirolimus was then dissolved in the chloroform/polymer mixture to give a final sirolimus/polymer ratio of 30% (w/w). The mixture was applied to an electro-grafted p-BuMA stent, mounted on a rotating mandrel, by spraying with a fine-meshed nozzle having the following parameters:

table 4: parameters of spraying

Two-layer coatings were applied on the luminal and abluminal (abluminal) faces of the stainless steel stent, which had a thicker (and better) abluminal surface than the luminal surface. Drying was carried out in a vacuum oven at 40 ℃. Using the parameters described above, the coating on the stent weighed 800+/-80 μ g, and the coating thickness was about 5 to 7 μm. The drug loading was 164+/-16 μ g.

Example 7: enhancement of adhesion of the electrografted p-BuMA underlayer to the depot polymer; and (4) a functional test.

Adhesion tests were performed to highlight the adhesion strength of the depot polymer layer on the pre-electro-grafted scaffolds. A two-layer coating on a stainless steel stent (18mm, Clearstream Technologies) was obtained according to examples 1 and 6.

This test was performed to simulate the erosion that the coating can experience during implantation. For this purpose, the coated stent is passed several times through a silicon tube simulating a coronary artery, after which the stent is deployed. After the test, the stent was examined by optical and scanning electron microscopy.

No coating delamination was observed for the substrate electro-graft scaffold: all 10 of the electrografted coated stents passed the simulated damage erosion test, while spray coated stents without the electrografted p-BuMA substrate showed severe delamination.

Example 8: in vitro drug Release study

In these embodiments, the time course of sirolimus release from the bilayer-coated stent was obtained according to the following experimental design:

an 18mm stainless steel stent was coated according to the experimental design given for the electro-grafting of p-BuMA in example 1 and for the biodegradable polymer spray coating in example 5. Each coated stent was immersed in a vial containing 1ml of release solution (99% phosphate buffer, 0.01M, pH 7.4/1% tween 20), stored at 37 ℃, and gently stirred (natural sizing). The release medium was removed periodically and replaced with a new one. The absorbance (arbitrary units) of the release medium was measured at a wavelength λ 278nm using a spectrophotometer Hitachi 3.

Sirolimus concentration was determined in triplicate using a calibration curve.

Fig. 1(a) and 1(B) illustrate the rapid release (a) and slow release (B) (cumulative release (%) versus time (days)) of sirolimus from the bilayer coating in vitro, respectively. For fast release the reservoir layer is a copolymer of lactide and glycolide, PLGA (120000g/mol) (50/50), and for slow release the biodegradable polymer is poly (lactide) (30000 g/mol).

The difference in drug release kinetics is directly related to the degradation rate of the biodegradable reservoir. Because 50/50PLGA polymer degrades faster than PLA.

The corresponding characteristics in vivo are shown in figure 2 (release rate versus time (days), which was obtained from the assay of the remaining drug on the grafted stent from NZ rabbits in the iliac-femoral model:

+ a feature of PLGA

x is characterized by the feature of PLA

In both cases, the released drug (sirolimus) and the loaded drug are the same. The figure shows that for fast release (PLGA) the drug is completely released and the release polymer disappears completely in 28 days, whereas for PLA the drug is only 60% released in 28 days, which is considered to disappear within two months.

Example 9: cytotoxicity Studies of electrotransport coatings

Potential cytotoxicity studies of the electro-grafted coatings were carried out according to standard ISO 10993-5.

The study was aimed at qualitatively and quantitatively evaluating the cytotoxicity of electroporated p-BuMA, as the extracts were tested after applying them to cells seeded in 96-well microplates.

Extraction was performed in triplicate in sterile, closed, chemically inert containers with medium containing fetal bovine serum (DMEM) at 37 ℃ over 96 hours. The ratio of the surface area of the electroporated p-BuMA to the volume of the extraction support was equal to 3cm²/ml。

The extract and its dilutions (50% and 10%) were placed on the cells and kept in contact for at least 24 hours. Cytotoxicity was measured with vital dye neutral red. The assay format maintained was common cell morphology (qualitative assessment) and the percentage of cell viability (quantitative assessment) was proportional to the number of viable cells (quantitative analysis), the percentage being based on absorbance obtained by reading at 540 nm.

Positive control: for each test, the test was carried out with a product providing a reproducible cytotoxic effect under test conditions: 3.2g/l phenol solution in culture medium (DMEM). If the percent mortality is about 100%, the test is satisfactory.

Negative control: under these test conditions, a control was performed with a substance (high density polyethylene) which produced no cytotoxic effect. The assay is satisfactory if the percent cell viability is 100%.

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Table 5: qualitative and quantitative evaluation of toxicity of electrotransport pBuMA substrates

The test performed on the coating extract of the electro-grafted substrate showed no evidence of cytotoxicity after twenty-four hours.

Example 10; a hemolysis study; direct contact testing.

In this example, hemolysis refers to the destruction of red blood cells in direct contact with the electrotransport coating. The hemolysis was studied at 1cm²According to ISO 10993-4 and ASTM F756-93, on p-BuMA and sterilized stainless steel specimens.

Repair of human blood matrix: citrate anticoagulated human blood was obtained in sterile conditions from three donors. Blood was used within 1 hour.

Dilution of blood matrix: hemoglobin concentration of each blood was evaluated and was 97.95. + -. 8.32-111.86. + -. 3.90-91.05. + -. 0.94 mg/ml.

The free plasma hemoglobin must be less than 1mg/ml (0.30-0.32-0.28 mg/ml).

The total hemoglobin content of each blood sample was adjusted to 25.01 + -2.5 mg/ml by dilution with an appropriate amount of physiological saline (25.66 + -0.05 mg/ml-26.19 + -1 mg/ml-25.37 + -0.69 mg/ml).

And (3) hemoglobin detection:

hemoglobin of blood: mu.l of blood was mixed with 5ml of Drabkin's reagent (Sigma-525-2) (15 min). The absorbance (arbitrary units) was measured with a spectrophotometer at λ 540 nm. Hemoglobin concentration was determined using a calibration curve from 0.036 to 0.72mg/ml prepared using a reference standard (hemoglobin standard, Sigma-525-18).

Plasma hemoglobin: in a hemolysis tube, 100. mu.l of plasma was mixed with 5ml of Drabkin's reagent (15 min). The absorbance (arbitrary units) against Drabkin's reagent was measured at a wavelength of 540nm using a spectrophotometer (Kontron). Hemoglobin concentrations were determined in triplicate using a calibration curve from 0.036 to 0.720mg/ml prepared using a reference standard (hemoglobin standard, Sigma-525-18).

Static test：

Under sterile conditions, 5ml of each blood substrate was delivered into a screw-cap tube containing the test substance. The ratio of the surface area of the test substance sample to the volume of the blood substrate was 3cm²And/ml. The positive control consisted of 200. mu.l of blood substrate plus 10ml of water.

The negative control consisted of blood matrix alone.

The tubes were capped and stored stationary in a suitable tube rack at 37 ℃ for 4 hours. At the end of the specified incubation time, all tubes were centrifuged (100XG, 15 min). The supernatant fraction of each cell-free plasma was transferred to a 15ml tube (polypropylene, sterile) and centrifuged (700XG, 5 min). The supernatant was carefully removed for subsequent hemoglobin analysis.

And (3) hemoglobin measurement: 1ml of the supernatant was mixed with 3ml of Drabkin's reagent. The absorbance was measured at λ 540 nm. Calibration curves from 0.03 to 0.72mg/ml were prepared using reference standards (hemoglobin standards, Sigma-525-18).

Plasma hemoglobin concentration in each supernatant was determined using a calibration curve.

The Hemolytic Index (HI) was calculated according to the following formula:

table 6 shows the hemoglobin levels in the supernatants and table 7 gives the corresponding Hemolytic Index (HI). Using 3 blood, the average HI for the negative controls obtained was 0.35 ± 0.04% evaluated in triplicate. The mean HI in the presence of p-BuMA substrate was 0.29. + -. 0.03% assessed in triplicate using 3 blood.

Table 6: hemoglobin level in supernatant

Table 7: index of hemolysis

The results show that the samples of the electro-grafted coating have no hemolytic properties when in direct contact.

Example 11: local tolerance of electro-graft scaffolds after implantation in rabbits

One objective of this study was to evaluate the local tolerance of the electro-grafted p-BuMA stent compared to a bare metal stent. Stents coated with an electro-graft coating (stainless steel, 18mm length) were coated according to the experimental design given in example 2, sterilized with ethylene oxide using an approved standard experimental design (43 ℃, 50% relative humidity).

Experimental procedure

1-site of implantation

Coated or uncoated stents were implanted into the right and/or left iliac artery sites of each animal for 4 weeks.

2-preparation and anaesthesia of animals

Pre-anesthesia administration of atropine (atropinum furifucum, Aguettat, France) to rabbits was performed according to the intramuscular route of the internal standard method using teletamine-zolazepam (Zoletil)100, VIRBAC, France) 25mg/kg and xylazine (Rompun)2% BAYER AG, Germany) 5mg/kg anesthesia. Cutting off the skin of the surgical site, and applying antiseptic soap (Vetedine)savon, VETOQUINOL, France) by washing with povidone iodine (Vetedine)Solution, VETOQUINOL, france) sterilized.

Prior to implantation of each stent, the following drugs were administered to the femoral artery via an introducer:

·Aspegic(synthalbapo, france), 50 mg.

·Heparine Choay(synthalabo, france), 50 IU.

Also, prior to each angiogram, the following vasodilating drugs were administered to the femoral artery:

Corvasal(linsidomine, 0.06mg, AVENTIS, France).

3-Pre-procedural angiography (Pre-procedure angiography)

One carotid artery was exposed and a 5 or 6Fr introducer cannula was introduced. 5 or 6 guide catheters and Guide Wires (GW) are advanced through the cannula to the distal aorta. By injecting contrast material (Hexabrix) with a Philips BV212 apparatus320, laboratories GUERBET, france) to perform angiographic mapping of the iliac vascular tree. The diameter of each artery is reported. The target overstretch after stent implantation is approximately 1.2.

4-Placement of the Stent

Stents were implanted into the iliac artery (1 or 2 stents per animal) according to the following procedure:

a Guide Catheter (GC) and GW are inserted into the target site.

Full retraction of the GW.

The stent deployment system is inserted into the target site.

Deployment of the stent is under a defined balloon pressure (═ 8 atmospheres).

Implantation of a stent in the common iliac artery.

The delivery system was withdrawn from the GC.

5-post procedural angiography

An immediate assessment of the patency of the implanted artery was performed using angiography. The diameter of each artery was reported and the resulting hyperstretch was calculated.

6-pharmacological treatment and Observation period

Animals were observed daily for any clinical abnormalities. Anticoagulant treatment was started the day before the implantation step, with each daily dose for 30 days: aspegic(aspirin 100mg/ml, SYNTHELABO, France), intramuscular, 50 mg/day.

7-sacrifice and sampling

By lethal injection of barbiturate (Dolethal)^NDLaboratoiiresVETOQUINOL, France) to sacrifice the animals. Visual inspection of the external surface of the implanted artery was performed: any criteria for local intolerance (inflammation, necrosis, bleeding or any other damage) were observed and reported. Macro photography is performed. Samples were identified and fixed in 10% buffered formaldehyde solution for histopathology.

Preparation of 8-histopathological samples

The implant site was dehydrated in increasing concentrations of alcohol solution and embedded in PMMA (polymethylmethacrylate). A distal portion was obtained by a Donath modified microhardness scoring and grinding technique (Donath K., Brunner G.: A method for the study of uncalcined bones and teeth with soft tissue attached, J.Oral.Pathol.11; 318-. The fractions were stained with a modified Paragon staining method for qualitative and quantitative analysis.

9-explanation

Histological sections were examined under light microscopy (NIKON Eclipse E600, fitted with x4, x10, x20 and x40 lenses, equipped with a digital camera DN 100 NIKON). Semi-quantitative histological evaluation was performed according to ISO 10993-6. Particular attention should be paid to the presence of fibrous tissue, fibrin, degenerative phenomena, necrosis, smooth muscle cells, elastic sheet swelling, ageing of inflammatory cells and substances and the appearance of thrombi.

A histological photomicrograph was performed. Each parameter was evaluated according to the following evaluation scale.

0: is absent from

1: limited

2: of moderate degree

3: is remarkable in

4: severe degree

These parameters allow accurate assessment of any inflammation, foreign body response and immune response. The formation of the inner membrane (Neointimal) was assessed quantitatively.

Results

1-histopathological analysis

Semi-quantitative analysis is reported in Table 2

2-general observations

The stent struts exhibit a square shape with rounded corners. No substantial change in the microtest was observed in the samples.

3-uncoated Stent (control product)

All stents are fully deployed, allowing good incorporation into the vessel wall. The scaffold framework was incorporated into medium thickness intimal tissue containing moderate numbers of smooth muscle cells, fibroblasts, and limited macrophage permeate. One sample (animal, n ° 3, right) showed limited elastic sheet breakage without inside protrusions. It was suspected that there was a limited amount of proteoglycan material present in the intimal tissue of one sample (animal, n ° 11, left). No thrombus was observed.

4-coated Supports (test products)

The thickness of the fibromuscular intima layer covering the stent frame is comparable to or slightly thinner than that of the reference group. These findings were obtained in control animals. Due to the limitations of the number of samples and the observation that can be evaluated, no conclusions were drawn about the biologically significant findings. The macrophagic (macrophagic) reaction had a slight magnification (slice) similar to the reference group. No thrombus was observed.

10 out of a total of 14 animals were successfully implanted with coated (test product) and/or uncoated stents (control product). In this study, the artery overstretching after stent implantation reached approximately 1.1 to 1.4 times the starting artery diameter. No macroscopic lesions (necrosis, inflammation, bleeding) were found in the samples obtained from 8 surviving animals (n ═ 6, uncoated stents; n ═ 7 coated stents) after 1 month of implantation. No signs of clogging were observed in the sacrificed animals.

Table 8: semi-quantitative histopathological analysis:

r is right; l is left; m is the average value

Results of rabbit histology part retrieval of scaffold 1 month after implantation: for all the experimental or control groups, there were no signs of local intolerance reactions, there was mild fibromyointimal proliferation in stenosis with comparable results.

Conclusion

The major histopathological findings were as follows:

all test and control stents were fully deployed, well binding to the vessel wall without thrombus.

One month after implantation, no signs of local intolerance were observed for all the test or control stent groups.

The test and control stent groups showed comparable results in terms of stenosis, with mild fibromuscular intimal proliferation.

Moreover, in the rabbit model (ISO 10993), the electro-grafted layer was able to prevent cracking and delamination of the biodegradable polymer layer and showed equal, if not better, re-grafting compared to the stainless steel bare metal stent.

Example 12: re-transplantation in p-BuMA electro-stented rabbits at 14 and 28 days, compared to BMS

The experimental design of example 3 was followed to coat a cobalt chromium stent with an approximately 200nm electro-grafted p-BuMA layer. Under general anesthesia, 20 stents (18mm, bare metal, n-10, and coated with an electro-grafted p-BuMA layer, n-10) were placed in the iliac femoral artery of ten new zealand white rabbits.

The first group of 5 animals was sacrificed on day 14, followed by the second group on day 28. According to Circulation in Finn et al,112the experimental design described in 270(2005) extracted the iliac femoral artery, prepared for longitudinal sectioning. The cross-section was examined by SEM and endothelial coverage was estimated from the SEM images (above). The results are summarized in tables 9 and 10 below:

table 9: endothelial coverage (%)

Table 10: endothelial coverage (%)

These results indicate that endothelial coverage (as calculated from SEM analysis of longitudinal sections) was higher or comparable on and between struts on stents coated with an electroporated p-BuMA layer compared to bare metal stents. It should be particularly noted that the re-implantation is effective as early as 14 days after implantation of the electro-graft stent, which indicates that the optimum according to this technique will benefit directly from this result and reduce the period of drug release to a minimum in order to promote a healing effect (pro-healing effect).

Example 13: local tolerability after complete coating of the scaffold in pigs

The pig test was carried out for 60 days with a combined layer prepared by electro-grafting a p-BuMA lower layer (150nm) coated with a biodegradable release layer (5 μm) of PLGA (polylactide-co-glycolide). Briefly, sixteen domesticated male pigs (25 to 30kg) underwent placement of 32 stents (18mm long, bare metal, n-16, and double coated stents, n-16) under general anesthesia on the left anterior descending artery (IVA) or the left coronary artery branch (Cx).

Segments with an average coronary diameter of 2.5mm were selected by angiography using a quantitative coronary with a stent to artery ratio of about 1.2. Next, a stent-secured balloon catheter is advanced over a standard guide wire to a preselected coronary section for deployment. The balloon catheter was inflated to 10 atmospheres at a time in 10 seconds and then slowly withdrawn leaving the stent in place (without prior or subsequent expansion).

Coronary IVUS

To assess the extent of intima formation in vivo, IVUS was performed 8 weeks after stent implantation.

Arterial samples:

the heart was excised 8 weeks after stent implantation. IVA, Cx and CD were removed, washed in Phosphate Buffered Saline (PBS), and then prepared for histomorphometry, immunochemical analysis or scanning electron microscopy.

Histomorphometry

The samples were fixed in formalin (3%) at 4 ℃ for 12 hours, dehydrated in a gradient ethanol series (70 ° to 100 °,4 ℃) and acetone for 24 hours, and then embedded in glycomethlymethacrylate (gma). For each sample, 50 μm thick sections were cut (Isomet, Buehler France) and stained with Verhoeff-van Giesen for analysis. The tissue sections were observed (Nikon E-600, Nikon, France), digitized, and morphometric (Metamorph, France). The thickened intima was quantified by morphometric analysis at 5 sections per arterial segment. The area of the inner membrane, e.g. the area from the in vivo elastic sheet (IEL) to the luminal edge, is determined, and the area of the medium is the area between the IEL and the outer elastic membrane. The thickened intima is expressed as the ratio [ (intima area/intima area + media area) ].

Immunochemical analysis

At the end of the drying process, the stent was removed, the artery embedded in a paraffin block, cut into 4 μm thick sections, and then immersed in 3% aqueous hydrogen peroxide (Sigma, france) to inhibit endogenous peroxidase activity. Nonspecific staining was blocked by incubation in 5% bovine albumin PBS for 10 min. After washing twice in PBS, sections were cultured in various antibodies (anti-MIB 1, alpha-actin, factor VIII, macrophages (AM-3K)). Two independent observers counted stained cells in the intima and media regions.

Scanning electron microscopy

For this purpose, the samples were fixed with 4% glutaraldehyde, 0.1M phosphate buffer, pH 7.2 for 1 hour at 4 ℃ and washed in PBS for 1 hour. Subsequently, they were dehydrated from CO by using a gradient of ethanol and pure acetone₂(CPD 010BAL-TEC AG, Liechtenstein) critical point. The coating samples were spray coated with Au/Pd (Emscope Ashford UK) for observation using scanning electron microscopy of secondary electrons (JSM 6300Jeol Tokyo Japan).

2.8. Statistical analysis

All experiments were performed in triplicate and the results are expressed as mean ± SD. Treatment with antiplatelet therapeutics (Plavix 300Mg and aspirin 75Mg) was started on the day before catheterization, continued to usual amounts (Plavix 75Mg and aspirin 75Mg per day) throughout the duration of the study (6 hours, 1 month and 2 months thereafter.) pigs were catheterized under fluoroscopy (Seldinger) by femoral approach probe "EBU" (Medtronic) was placed at the left coronary trunk mouth, with selective opacification of the coronary reticular structure, after injection of 50UI/kg heparin, initial intracoronary echographic control (IVUS) (atlanis Plus 40mhz, Boston) was performed, initial IVUS made it possible to estimate coronary artery diameter and guide stent implantation to obtain a stent/artery ratio of 120%. next, a stent was placed in the middle section of the vessel (12 atmospheres, dryness of 10). After a new coronary (corographic) control and IVUS to protect the stent placement, the material was removed and maneuvers were applied to press at the site of the femoral puncture until hesmostatis was obtained. After a two month life span, a new catheterization was performed as described above for coronary control. Evaluation of in-stent stenosis and intimal proliferation was performed by new IVUS.

The results of this study will first show that the biodegradable release layer disappears after the first 4 weeks, thereby releasing 100% of the drug. Indeed, one can only see by SEM that a "rough" electroformed p-BuMA layer is characterized by its "lunar" (reproducible "crater" surface irregularities; although it is homogeneous and contains polymer everywhere, even in "lunar pores"). By observing the surface of the stent by SEM 30 days after implantation, one can see the "crescent-shaped crater" characteristic of the electro-grafted p-BuMA, confirming that the biodegradable layer has completely disappeared and, therefore, the drug is completely released. The disappearance of the biodegradable layer was further confirmed by ToF-SIMS analysis of the surface and inside of the artery as described above, which showed the absence of the biodegradable polymer drug.

Given the observed re-transplantation at week 8, electro-transplantation of p-BuMA tended to be by appropriate re-transplantation of endothelial cells.

IVUS results demonstrate very good tolerance of the bilayer coated scaffold, since after 8 weeks of implantation, very little intimal proliferation and very little smooth muscle cell proliferation were observed, which can be confirmed by immunohistological studies, which confirm that the coating is very safe without inflammation, as shown by HES staining, which is a complete endothelylation von- (willebrand staining).

Example 14: "Low pressure" spray device for preparing DES with good bonding surface to electro-grafted layer

The apparatus consists of a glove box with a transparent wall. The X-Y scanning system is placed on top of the outside of the cabinet and is moved by an outside magnet, which is further operated by another magnet on the ceiling inside the cabinet for X-Y movement. The latter internal magnet may be further connected to the nozzle.

The remote wall of the tank has a male appendage connected to an external electric motor, which can be made to rotate at a controlled speed by adjusting the voltage of said motor from the front panel of the device.

The holder was placed on the probe, which was further plugged with a tip on the sample holder. These tips can rotate around the entire circumference relative to the sample holder: they are attached inside the usual rotating rods of the sample holders, the ends of which-behind the sample holders-have female appendages that can be plugged with male appendages in the distal wall (far wall) of the magazine. Thus, when all the supports are placed on the probe, where they can themselves jam the tip, and when the sample holder is inserted on the distal wall of the case, all the supports rotate simultaneously and at the same adjustable speed.

The X-Y system is controlled by an external computer and pushes the sequence of movements and sprays of nozzles located at the top of each rack, starting the sprays one after the other as the entire length of rack is moved forward in one direction and in the opposite direction, before the spray ends, will startMove to another carriage and begin spraying again. The sample holder can hold 20 per batch, therefore, the nozzle is transferred from the back of rack #20 to rack #1 for a second spray: in the entire length of the sample holder, the spray of each rack is essentially not "visible" during a period corresponding to 19 times the Ts time of spraying from one rack + the time T of cleaning the nozzle₀1 time of. Thus, it can be seen that all stents had identical experimental designs and the coating apparatus was confirmed to be highly reproducible.

The quality detection error results (coating on a series of 53 DBS qualities which have been sprayed using equipment as described above) showed an acceptance criterion of 15% relative to the target quality (and hence target drug dose), with only 3 DES being rejected as out of specification, yielding an overall yield of 94.2%. Even with the more stringent 10% tolerance imposed, the system gives a yield of 86.5%, which is substantially higher than most present in industrial systems (since its usual specification for drug dosage is about 20%).

We attribute this property to the very high reproducibility of wet/low pressure systems, which is ideally facilitated by the wettability provided by a suitable electron transfer layer. We also feel that the relationship between the drug concentration in the spray solution and the last polymer layer is highly linear and highly reproducible for low pressure, gas driven, nozzles, even though there is no strict one-to-one correspondence.

Claims (14)

1. A drug eluting stent, comprising:

a stent frame;

an electro-graft coating applied to the stent framework, and

a biodegradable polymer coating containing a drug, coated on the electro-graft coating by interdigitation,

wherein the electro-implant coating:

thicker than 100 nm;

has the same wettability as the biodegradable polymer coating;

has a glass transition temperature less than that of the biodegradable polymer coating for thermal cross-bonding; or

When the biodegradable polymer coating is applied by coating, the electrographic coating is swollen with at least a solvent capable of dissolving the biodegradable polymer or a solvent comprising a dispersion of the biodegradable polymer or a component thereof;

the electrographic coating is prepared from a monomer selected from the group consisting of a vinyl compound, an epoxide, and an aryl diazonium salt.

2. The drug eluting stent of claim 1, wherein the stent frame comprises a metal matrix.

3. The drug eluting stent of claim 1 or 2, wherein the stent framework comprises a material selected from the group consisting of: stainless steel, nickel, tantalum, cobalt-chromium MP35N or MP20N alloy, platinum, titanium, and combinations thereof.

4. The drug-eluting stent of claim 1, wherein the electro-graft coating has a thickness of 100nm to 1.0 micron.

5. The drug-eluting stent of claim 1, wherein the monomer is selected from the group consisting of butyl methacrylate, methyl methacrylate, hydroxyethyl methacrylate, epsilon caprolactone, and 4-aminophenyldiazotetrafluoroborate.

6. The drug-eluting stent of claim 1, wherein the drug-containing biodegradable polymer coating applied over the electro-graft coating contains a bioactive agent.

7. The drug eluting stent of claim 6, wherein the bioactive agent is selected from the group consisting of antisense agents, anti-tumor agents, antiproliferative agents, antithrombotic agents, anticoagulants, antiplatelet agents, antibiotics, anti-inflammatory agents, gene therapy agents, recombinant DNA products, recombinant RNA products, collagen derivatives, proteins, carbohydrates, and carbohydrate derivatives.

8. The drug-eluting stent of claim 1, wherein the drug-containing biodegradable polymer coating is selected from one or more biodegradable polymers.

9. The drug eluting stent of claim 8, wherein the biodegradable polymer is selected from the group consisting of biodegradable copolymers and biodegradable block polymers.

10. The drug-eluting stent of claim 9, wherein the biodegradable polymer is selected from the group consisting of polyglycolide, polylactide, polycaprolactone, polyglycerol sebacate, polycarbonate, biopolyester, polyethylene oxide, polybutylene terephthalate, polydioxanone, collagen matrix with growth regulators, proteoglycans, glycosaminoglycans, chitin, dextran, and mixtures thereof.

11. The drug-eluting stent of claim 1, wherein the drug-containing biodegradable polymer coating has a thickness of 1 to 200 microns.

12. The drug eluting stent of claim 1, further comprising a biodegradable topcoat.

13. A drug eluting stent according to claim 12, wherein the topcoat layer is prepared using the same composition as the biodegradable coating release layer.

14. The drug-eluting stent of claim 8, wherein the biodegradable polymer is selected from the group consisting of poly- β -hydroxyalkanoates.