HK1150229B - Drug-delivery endovascular stent and method for treating restenosis - Google Patents

Description

Drug-releasing intravascular stent and method for treating restenosis

The present application is a divisional application of international application with application number PCT/US2003/012750 entitled "drug-releasing intravascular stent and method for treating restenosis" filed 24/4/2003, which entered the chinese national phase at 25/10/2004, application number 03809311.1.

The invention belongs to the field of the following:

the present invention relates to a drug-releasing intravascular stent and a method for treating restenosis.

Background

A stent is an endovascular implant, generally tubular in shape, generally having a reticulated, interconnected wire tubular structure that is expandable, may be permanently placed in a blood vessel to provide mechanical support to the vessel and maintain or reestablish a blood flow passageway during or after angioplasty. The support structure of the stent is designed to prevent early collapse of a weakened or damaged vessel caused by angioplasty. The placement of the stent has been shown to prevent negative remodeling and spasm of the vessel during the months of the healing process of the damaged vessel wall.

During the healing process, inflammation caused by angioplasty and stent placement injury often results in smooth muscle cell proliferation and regrowth within the stent, thus occluding portions of the blood flow passageway, thereby reducing or eliminating the beneficial effects of angioplasty/stenting procedures. This process is called restenosis. Because of the thrombogenic nature of the stent surface, blood clots may also develop within the newly placed stent, even with stents made of biocompatible materials.

Since the current practice is to inject a strong antiplatelet agent in the blood circulation during angioplasty, there is no large blood clot formation during and shortly after surgery, but some thrombus is always present, at least visible on the stent surface under the microscope, which is thought to play an important role in the early stages of restenosis by establishing a biocompatible matrix on the stent surface, on which smooth muscle cells subsequently attach and proliferate.

Stent coatings containing bioactive agents are known, and are designed to reduce or eliminate thrombosis or restenosis. The bioactive agent is dispersed or dissolved in a biodurable or biodegradable polymer matrix attached to the surface of the stent filaments prior to stent placement. Following stent placement, the bioactive agent diffuses out of the polymer matrix, preferably into the surrounding tissue, for at least 4 weeks, in some cases up to a year or more, ideally matching the time course of restenosis, smooth muscle cell proliferation, thrombosis, or a combination thereof.

If the polymer is biodegradable, in addition to releasing the drug by a diffusion process, the bioactive agent may also be released as the polymer is degraded or dissolved, making the bioactive agent more accessible to the surrounding tissue. Biodegradable and biodurable stents are known, the outer surface of which, or even the entire polymeric material, is porous. Such a stent is disclosed, for example, in PCT publication No. WO99/07308 (of the same applicant as the present application), which is incorporated herein by reference. When biodegradable polymers are used as coatings for drug delivery, numerous patents describe that porosity can aid tissue ingrowth, make polymer degradation more predictable, or modulate or increase the rate of drug release, such as US patents: 6,099,562, 5,873,904, 5,342,348, 5,873,904, 5,707,385, 5,824,048, 5,527,337, 5,306,286, and 6,013,853.

Heparin and other anti-platelet or anti-thrombolytic surface coatings are known, which are chemically bonded to the surface of the stent to reduce thrombosis. Heparinized surfaces are known to interfere with the human coagulation cascade, preventing the attachment of platelets (the precursor of thrombin) to the surface of the scaffold. Stents with heparin on the surface and active agent stored within the coating have been described (see, e.g., U.S. Pat. Nos. 6,231,600 and 5,288,711).

There are many active agents that inhibit smooth muscle cell proliferation and thus inhibit restenosis, and these agents have been proposed for release from intravascular stents. For example, U.S. Pat. No. 6,159,488 describes the use of quinazolinone derivatives; us patent 6,171,609 describes the use of Taxol; U.S. Pat. No. 5,176,98 describes the use of paclitaxel, a cytotoxic agent that is considered to be an active ingredient in the drug Taxol. Metallic silver is mentioned in us patent 5,873,904. Tranilast, a membrane stabilizer believed to have anti-inflammatory properties, is disclosed in us patent 5,733,327.

More recently, rapamycin, an immunosuppressant reported to inhibit smooth muscle cell and endothelial cell growth, has been shown to have improved efficacy against restenosis when administered from a polymer coating on a stent. See, for example, U.S. Pat. nos. 5,288,711 and 6,153,252. In addition, PCT publication No. WO97/35575 proposes the use of the macrocyclic triene immunosuppressive compound everolimus and related compounds for the treatment of restenosis by systemic administration.

Ideally, a compound selected to inhibit restenosis by drug release from a stent should possess the following three characteristics:

1. due to the low profile of the stent, i.e., thin polymer matrix, it is desirable that the compound released from the thin polymer coating be sufficiently active to produce a sustained therapeutic dose for a minimum of 4-8 weeks.

2. The compound can effectively inhibit smooth muscle cell proliferation under the condition of low dose.

3. The endothelial cells on the luminal surface of the blood vessel are typically damaged during angioplasty and/or stenting, and the compound should allow for re-growth of endothelial cells within the vessel lumen, provide for return of vascular homeostasis, and promote normal and important interactions of the vessel wall and blood flow through the vessel.

Summary of The Invention

In one aspect, the invention includes an intravascular stent for placement at a site of vascular injury to inhibit restenosis at the site. The stent is composed of a structural member or structure formed of one or more filaments and has a biodegradable drug-releasing coating of thickness 3-15 microns on the stent body filaments, the coating being composed of (i) 20-60% by weight of a poly-dl-lactide polymer substrate and 40-80% by weight of an anti-restenotic compound. The polymer base layer has a thickness of 1-5 microns and is positioned between the stent body filaments and the coating and functions to stabilize the coating on the stent filaments. The stent is expandable from a contracted state in which the stent may be delivered to a site of vascular injury via a catheter, and an expanded state in which the stent coating is in contact with the site of vascular injury. The stent coating is effective to release a restenosis-inhibiting amount of the compound for a period of at least 4 weeks after the stent is placed in the site of vascular injury.

In various exemplary embodiments, the anti-restenotic compound is a macrocyclic triene immunosuppressive compound, the stent body is a wire structure, and the bottom layer immediately adjacent to the wire is formed of parylene polymer and has a thickness of 0.5 to 5 microns; the thickness of the coating is 2-10 microns. The compound may be present in the coating in an amount of 50-75% by weight.

Exemplary macrocyclic triene immunosuppressive compounds have the following structure:

wherein (i) when R 'is H (R' replaces H at position 28O), R is H or CH₂-X-OH and X is a linear or branched alkyl group containing from 1 to 7 carbon atoms, or (ii) at least one of R' and R is of the form:

m is an integer of 1 to 3, R₁And R₂Are each hydrogen or alkyl having 1 to 3 carbon atoms, or R₁And R₂Together with the nitrogen atom to which they are attached form a saturated heterocyclic ring having 4 carbon atoms. In one exemplary compound, everolimus, R' is H and X is-CH₂。

According to another aspect of the present invention, the above-described stent is used in a method for inhibiting restenosis at a site of vascular injury. In this method, a stent is introduced into a vascular injury site and then expanded to bring the stent coating into contact with the vascular injury site. The coating is effective to release a restenosis-inhibiting amount of the compound for a period of at least 4 weeks.

In another aspect, the invention includes an intravascular stent for placement at a site of vascular injury to inhibit restenosis at the site. The stent is composed of a structural member or structure formed by one or more filaments, and a drug-releasable coating with a thickness of 3-25 microns is provided on the filaments of the stent body, the coating is composed of (i) 20-70% by weight of a polymer substrate and (ii) 30-80% by weight of a macrocyclic triene immunosuppressive compound, and the molecular structure is as follows:

wherein R is CH₂-X-OH, X being a linear group containing from 1 to 7 carbon atoms.

The stent is expandable from a contracted state in which the stent may be delivered to a site of vascular injury via a catheter, and an expanded state in which the stent coating is in contact with the site of vascular injury. The stent coating is effective to release a restenosis-inhibiting amount of the compound for a period of at least 4 weeks after the stent is placed in the site of vascular injury.

In various exemplary embodiments, R is CH₂-X-OH, X being-CH₂The stent body is a wire structure and the polymer substrate in the coating is polymethyl methacrylate, ethylene vinyl alcohol copolymer or poly-dl-lactide polymer.

In an exemplary embodiment, the polymer substrate in the coating is formed from biodegradable poly-dl-lactide having a thickness of 3 to 20 microns; the initial concentration of the compound contained in the coating is 20-70% of the total weight of the coating. In particular, when the compound content of the coating is greater than 40% by weight, the stent may also have a base layer of parylene polymer between the filaments of the stent body and the poly-dl-lactide coating substrate, with a thickness of 1-5 microns.

Alternatively, both the stent body and the coating substrate may be formed from biodegradable polymers, such as poly-l-lactide or poly-dl-lactide forming the stent body filaments and poly-dl-lactide forming the coating substrate.

When the stent is placed in a desired position, the stent coating is exposed to blood flow through the stent in the expanded state. In this embodiment, the coating may further contain a bioactive agent such as an antiplatelet agent, a fibrinolytic agent, or a thrombolytic agent in soluble crystalline form. Examples of antiplatelet agents, fibrinolytic agents or thrombolytic agents are heparin, aspirin, hirudin, ticlopidine, eptifibatide, urokinase, streptokinase, Tissue Plasminogen Activator (TPA) or mixtures thereof.

In yet another aspect, the present invention provides an improved method for treating restenosis at a site of vascular injury by placement of an intravascular stent designed to release a macrocyclic triene immunosuppressive compound over an extended period of time. The improvement comprises the use of a macrocyclic triene immunosuppressive compound having the structural formula:

r is CH₂-X-OH, X being a straight chain alkyl group containing 1 to 7 carbon atoms. In an exemplary compound, X is-CH₂-。

Various exemplary embodiments of the stent composition are described above.

A novel method of applying a drug-containing polymer coating to stent body filaments is also disclosed. The method uses an automatic controller to regulate the flow of polymer or drug-containing polymer solution onto the stent body filaments to achieve one of a variety of stent coating characteristics, including a uniform coating thickness on one or more sides of the stent body filaments; the thickness of the coating on the external (or internal) surface of the stent is thicker than that on the other surface; the inner and outer coatings contain different drugs; and/or a graded or discontinuous coating may be applied to the stent body.

These and other objects and features of the present invention will become more apparent from the following detailed description of the invention when taken in conjunction with the accompanying drawings.

Brief description of the drawings

Fig.1 and 2 illustrate an intravascular stent having a wire body formed in accordance with an embodiment of the present invention, fig.1 being a contracted state of the stent; fig. 2 is an expanded state of the stent.

Fig. 3 is an enlarged cross-sectional view of the coated wire in the stent of fig. 1.

FIG. 4 is an enlarged cross-sectional view of a coated degradable polymer stent.

Fig. 5A and 5B are schematic illustrations of a polymer coating process suitable for producing a coated stent of the present invention.

Fig. 6 is a schematic view of a biodegradable polymer stent constructed in accordance with the present invention mounted on a catheter for delivery to a vascular site.

FIGS. 7A and 7B are graphs of the release of everolimus from a stent of the present invention.

Fig. 8 is a cross-sectional view of the stent of the present invention after placement in a vascular site.

FIGS. 9A-9C are histological sections of blood vessels after 28 days of insertion of a bare metal stent.

FIGS. 10A-10C are histological sections of blood vessels after 28 days of implantation of a wire stent having a polymer coating.

FIGS. 11A-11C and 12A-12C are histological sections of blood vessels after 28 days of implantation of a polymer-coated wire stent containing everolimus.

FIG.13 is an enlarged histological section of blood vessels, in which the filaments of the stent used in FIGS. 11A-11C are seen, which filaments have been overgrown with new tissue to form a healed vessel wall.

FIG. 14 is a plot of area of stenosis as a function of lesion size 28 days after implantation of various stents, including stents constructed in accordance with the present invention.

FIG. 15 shows a plot of the correction function, lesion on the Y-axis and B/A (balloon/artery) ratio for stent placement on the X-axis.

Detailed Description

First section, intravascular stent

Fig.1 and 2 show a stent 20 constructed in accordance with the present invention in a contracted and expanded state, respectively. The stent includes a structural member or structure 22 and an outer coating for containing and releasing an anti-restenosis compound, as further described below with reference to fig. 3 and 4.

A、Support body

In the illustrated embodiment, the stent body is comprised of a plurality of tubular members, such as members 24 and 26, connected by filaments. Each member has an expandable zigzag or sinusoidal wave configuration. The bodies are connected by axial connectors, such as connectors 28, 30 connecting the peaks and troughs of adjacent sections. It will be appreciated that this configuration allows the stent to expand from a contracted state (as shown in fig. 1) to an expanded state (as shown in fig. 2) with the length of the stent unchanged or slightly changed. At the same time, the relatively few connections between the peaks and valleys of adjacent tubular members allows the stent to flex. This property is particularly important for stents that are introduced into a vascular site in or on a catheter in a contracted state. The stent has a typical contracted state diameter of 0.5-2mm (FIG. 1), more preferably 0.71-1.65mm, and a length of 5-100 mm. In its expanded state (see fig. 2), the stent has a diameter at least 2 times, and even 8-9 times, that of its contracted state. Thus, a stent having a contracted diameter of 0.7-1.5mm can be radially expanded to a selected expanded state having a diameter of 2-8mm or greater.

Stents having such a stent body comprised of connected, expandable tubular members are known, as described in PCT publication No. WO99/07308, which is owned by the same applicant as the present application and incorporated herein by reference. Other examples are described in U.S. Pat. Nos. 6,190,406, 6,042,606, 5,860,999, 6,129,755 or 5,902,317, which are incorporated herein by reference. Alternatively, the structural member of the stent may be a continuous helical ribbon structure, i.e., the stent body is formed from a single continuous ribbon-like helix. The basic requirements of the stent body are that it be expandable when placed in a vascular injury site and that it be coated on its outer surface with a drug-containing coating that delivers the drug contained in the coating to the vessel wall (e.g., the media, adventitia, and endothelial layers of tissue) lining the target site in the vessel. Preferably, the stent body also has a mesh or open structure that allows endothelial cells to grow through the stent from the outside in.

B、Stent coating

According to an important feature of the invention, the stent filaments are coated with a drug-releasable coating consisting of a polymeric matrix and an anti-restenosis compound (active compound) distributed within the matrix for releasing the drug from the stent over a period of at least several weeks, typically 4-8 weeks, sometimes for 2-3 months or more.

Fig. 3 shows, in an enlarged cross-sectional view, a stent wire 24 having a coating 32 that completely covers all sides of the wire, i.e., the top (the side of the wire that forms the outer surface of the stent body), the bottom (the side of the wire that forms the inner surface of the stent body), and the opposite side of the wire. As will be discussed further below, the thickness of the coating is typically 3-30 microns, depending on the nature of the polymer matrix material comprising the coating and the relative amounts of polymer matrix and active compound. It is desirable that the coating be as thin as possible, such as 15 microns or less, to minimize the profile of the stent at the site of vascular injury.

The thickness of the upper (outer) surface coating should also be relatively uniform to promote uniform distribution of the released drug at the target site. Methods for producing a coating of relatively uniform thickness on a stent wire are discussed in the second section.

Figure 3 shows a polymer base layer 34 between the stent filaments and the coating. The purpose of the base layer is to aid in the adhesion between the filaments of the stent body and the coating, i.e., to stabilize the coating on the filaments. As will be seen below, this function is particularly valuable when the coating-forming polymer substrate contains a high percentage of anti-restenosis compound (e.g.35-80% by weight of compound). An example of an underlayer polymer is parylene, which is used to attach a polymer substrate formed of biodegradable poly-dl-lactide. Other suitable polymeric substrates are ethylene vinyl alcohol copolymers (EVOH),paryLAST^TMParylene, polysiloxane, TEFLON^TMAnd other fluoropolymers that may be deposited on the surface of the metal stent by plasma coating, other coating or precipitation processes. The thickness of the bottom layer is typically 1-5 microns.

The polymer forming the substrate may be any biocompatible polymeric material in which the contained compounds can be released by diffusion and/or by degradation of the polymeric matrix. Two well-known non-degradable polymers for coating substrates are polymethyl methacrylate, ethylene vinyl alcohol copolymer. Methods for preparing these polymers suitable for use in the stent body are described in US2001/0027340A1 and WO00/145763, both of which are incorporated herein by reference. Typically, the drug loading limit for the polymer is about 20-40% by weight.

Biodegradable polymers, especially poly-dl-lactide, are also suitable as coating substrate materials. In one embodiment of the invention, the coating is a biodegradable poly-dl-lactide polymer substrate, i.e. a poly-dl-lactic acid polymer, which can contain up to 80% (by dry weight) of the active compound dispersed within the polymer substrate. More generally, the coating comprises 35-80% by dry weight of active compound and 20-65% by dry weight of polymer. Exemplary coatings contain 25-50% by dry weight of the polymer matrix and 50-75% by weight of the active compound. For a detailed description of the synthesis of polymers and drugs for deposition onto stent filaments, see section two.

A variety of anti-restenotic compounds may be used in this embodiment, including antiproliferative agents such as TAXOL, antisense compounds, doxorubicin and most particularly macrocyclic triene immunosuppressive compounds having the general formula shown below. The latter class of compounds and their synthesis are described in U.S. Pat. nos. 4,650,803, 5,288,711, 5,516,781, 5,665,772 and 6,153,252; PCT publication No. WO97/35575, U.S. patent application Nos. 6273913B1, 60/176086, 20000212/17, and 2001002935/A1, all of which are incorporated herein by reference. Exemplary macrocyclic triene immunosuppressive compounds have the following structure:

wherein (i) when R 'is H (R' replaces H at position 28O), R is H or CH₂-X-OH and X is a linear or branched alkyl group containing from 1 to 7 carbon atoms, or (ii) at least one of R' and R is of the form:

m is an integer of 1 to 3, R₁And R₂Are each hydrogen or alkyl having 1 to 3 carbon atoms, or R₁And R₂Together with the nitrogen atom to which they are attached form a saturated heterocyclic ring having 4 carbon atoms. In one exemplary compound, everolimus, R' is H and X is-CH₂。

One preferred coating is formed from 25-50% by weight of a poly-dl-lactide polymer substrate and 50-75% by weight of a macrocyclic triene immunosuppressive compound, the coating thickness being 3-15 microns. The bottom layer is poly-p-phenylene dimethyl with a thickness of 1-5 microns. In this embodiment, the amount of compound is equivalent to 15 μ g drug/mm stent length.

In another embodiment, the coating is formed from 15 to 35 weight percent of a degradable or non-degradable polymeric substrate and 65 to 85 weight percent of a macrocyclic triene compound. The coating thickness is preferably 10-30 microns and the stent may also contain a 1-5 micron polymer base layer, such as a parylene base layer. In this embodiment, the amount of compound is equivalent to about 15 μ g of drug per mm of stent length. The active compound has the following structure:

wherein R is CH₂-X-OH, X being a straight chain alkyl group containing 1 to 7 carbon atoms. A preferred compound is everolimus, wherein X ═ CH₂. Compounds in which X is an alkyl group of 2,3, 4, 5,6 or 7 carbon atoms, whether used alone or in any combination, as well as the use of the foregoing compounds, including acetates of everolimus, are also suitable for use in the present invention.

The coating may also contain a second bioactive agent that can minimize or improve healing of blood-related events, such as clotting that may be stimulated by the presence of a native vascular injury or stent. Examples of the second active agent include antiplatelet agents, fibrinolytic agents, or thrombolytic agents in soluble crystalline form, or NO donors that stimulate endothelial cell healing and control smooth muscle cell growth. Examples of antiplatelet agents, fibrinolytic agents or thrombolytic agents are heparin, aspirin, hirudin, ticlopidine, eptifibatide, urokinase, streptokinase, Tissue Plasminogen Activator (TPA) or mixtures thereof. The amount of the second active agent in the stent coating is determined by the length of the period of time in which the active agent is required to provide therapeutic benefit. Typically, the active agent acts in the first few days after vascular injury and stent placement, although for some active agents a longer release period is required.

The second active agent may be included in the coating formulation and applied to the stent filaments according to known methods.

C、Biodegradable stent

In another embodiment, both the stent body and the polymer coating are comprised of biodegradable polymers, and the stent is completely resorbed over time. The stent is preferably an expandable helical stent having a stent body (not shown) formed from helical ribbon filaments. Self-expanding helical stents for implantation into blood vessels are described in U.S. Pat. No. 4,990,155, which is incorporated herein by reference.

Helical stents may be prepared using a preformed stent having a final expanded diameter slightly larger than the size of the lumen of the vessel to be treated with the helical stent (3.5 mm Outer Diameter (OD) ± 1mm is common for coronary arteries). Generally, the scaffold can be prepared by: and (3) casting the stent in the expanded state, and screwing the stent to be in a contracted state by taking the long shaft of the stent as an axis or radially pressing the stent to be in the contracted state so as to be arranged on the tip of the catheter and sent to the blood vessel. The total thickness of the stent is preferably 100-1000 microns, and the total length is 0.4-10 cm. In fact, an important advantage of this type of biodegradable stents is that relatively long stents, such as stents in excess of 3cm in length, can be easily delivered and placed at the site of vascular injury.

Methods for forming expandable ballstents using braided filaments of biodegradable polymers such as poly-l-lactide have been reported (U.S. patent 6,080,177). There is also a form of device that has been used to release drugs (us 5,733,327).

One preferred polymeric material for forming the stent is poly-l-or poly-dl-lactide (U.S. patent 6,080,177). As described above, the stent body and coating may be integral, forming an expandable silk stent containing an anti-restenosis compound throughout the stent. Alternatively, the biodegradable coating may be applied to a preformed biodegradable body, as described in more detail in the second section below. In the latter case, the stent body may be formed from one biodegradable polymer, such as poly-l-lactide, and the coating layer may be formed from a second polymer, such as poly-dl-lactide polymer. The coating, if applied to a preformed stent, has substantially the same composition and thickness characteristics as described above.

Fig. 4 shows a cross-sectional view of a filament, e.g. a helical ribbon, in the biodegradable stent just described, with the stent body and coating formed separately. There is shown an inner biodegradable stent filament 36 coated on all sides with a biodegradable coating 38. An exemplary coating is formed from poly-dl-lactide and contains 20-40% by weight of an anti-restenosis drug, such as a macrocyclic triene immunosuppressive compound, and 60-80% by weight of a polymeric substrate. In another embodiment, the coating comprises 45 to 75% by weight of the compound and 25 to 55% by weight of the polymer matrix. Other classes of anti-restenosis compounds, such as those described above, may be used in both of the above embodiments.

Biodegradable stents have the unique advantage of treating the entire blood vessel with a device that can be used either in conjunction with balloon angioplasty to pre-dilate the vessel in the presence of a large occlusion or as a prophylactic implant in patients at high risk of future potential occlusions. Because the stent is fully biodegradable, it does not interfere with the patient's subsequent opportunity to perform uncomplicated surgery on the blood vessels, as does an "all-metal coat," i.e., a string of metal-based drug eluting stents.

As noted above, a second active agent can be incorporated into the coating for release from the coating within a desired period of time after implantation. Alternatively, if a second active agent is used, the second active agent can be incorporated into the stent body filaments if the coating applied to the stent body does not cover the inner surface of the stent body. The coating methods described in the second section below with respect to the metal wire stent body are also applicable to polymer wire stent bodies.

A second part: stent coating method

Referring in more detail to the drawings, FIGS. 5A and 5B are schematic views of the stent coating process of the present invention. The polymer is dissolved in a compatible solvent to provide a polymer solution 40. At least one anti-restenosis compound and, if desired, a second active agent, in the form of a suspension or solution, are added to the polymer solution with the same solvent or different solvents. The entire mixture is placed in a pressurized reservoir 42. A liquid pressurizing pump 44 is connected to the reservoir.

Any pressure source can be used as the pressurizing pump as long as the mixed solvent can be pushed through the solution transfer pipe 46 at a designed speed. The booster pump 44 is controlled by a microcontroller (not shown) as is well known in the art of precision feed systems. For example: such a microcontroller may contain 4-Axis automatic feed devices (4-Axis Dispensing Robot, model I & J500-R and I & J750-R, available from I & J Fisnar, FairLawn, NJ, which may be controlled by a personal computer via an RS-232C exchange interface), or precision feed systems such as Automove A-400, available from Asymtek, Calcey, Canada. A suitable software program for controlling the RS232C interface may be available from including the Fluidmove system, which may also be available from Asymtek, Calcey, Canada.

Attached to reservoir 42, for example to the bottom of the reservoir, is a solution conduit 48 which delivers the solvent mixture to the surface of the stent. The pressurizable reservoir 42 and conduit 48 are mounted on a movable support (not shown) which allows the solvent conduit to be moved in small steps, for example 0.2mm each, or continuously along the long axis of the support (as indicated by the arrow X1). The movable carriage of pressurizing reservoir 42 and catheter 46 can also move the tip (distal end) of the catheter in small steps toward or away from the surface of the filament in the direction indicated by Y1.

The uncoated stent is clamped in a rotating clamp having at least one end in contact with the inner surface of the stent. The clip may be rotated axially by a stepper motor at a slight angle, such as 0.5 degrees at a time, so that the outermost surface of the stent structure may be coated with a catheter, as is well known in the art. The support may also be continuously rotated if desired. Methods of precisely positioning low volume liquid conducting devices are well known in the art of X-Y-Z solvent feed systems and may be used in the present invention.

The activation of the fluid pressurizing pump, the positioning of the fluid conduits X1 and Y1, and the positioning of the stent R1 are all controlled by a digital controller and computer software so that a metered amount of solution is applied precisely where needed to the stent surface, and then the solvent can be evaporated leaving a hardened polymer and active agent coating on the stent surface. The viscosity of the solvent mixture typically varies with the amount of solvent, which may range from 2 to 2000 centipoise, and may typically be 300 and 700 centipoise. Alternatively, the catheter may be fixed in a fixed position and the stent may be moved along its long axis in addition to rotation to complete the coating process.

The X-Y-Z positioning stage and the movable carriage are commercially available from I & J Fisnar. The solution conduit is preferably an 18-28 gauge stainless steel syringe with a lockable connector attached. Such a catheter is available from EFD corporation of East Providence, RI. See the guide for needle selection for the particular use of EFD. The preferred injection needles are "Poorless blunted stainless steel needles, Length 1/4" quick point-to-point injection of granular or thick material "re-ordered from 5118-1/4-B to 5121-1/4-B; reorder "oval stainless steel needle 51150VAL-B, coat thick slurry, sealant and epoxy on flat strip deposit"; reordering "anti-cyanoacrylate" from 5121-TLC-B to 5125-TLC-B provides additional deposition control for low viscosity liquids. Embossed and teflon lined ". A disposable pressurized solution reservoir is also available from EFD under stock numbers 1000Y5148 to 1000Y 5152F. Another needle useful in the present invention is a glass capillary tube having an inner diameter of about 0.0005 to 0.002 inches, such as about 0.001 inches, available in the VWR catalog under catalog number 15401 and 560 "Microhematocrit solvent tube" 60mm in length and 0.5 to 0.6mm in inner diameter.

Under the bunsen burner, the tube is further stretched to obtain the desired dimensions for precise coating of the polymer/drug/solvent mixture. Programmable microcontrollers and XYZ stages for operating stepper motors are available from Asymtek corporation. The invention also includes the use of more than one liquid feed tube to form the coating together, or in the same coating process, more than one removable solution reservoir equipped with different needles or different chemical agents containing solutions of different viscosities or consisting of multiple solutions. The cardholder and stepper motor system is commercially available from Edmund Scientific, n.j, Barrigton.

As noted above, the coating is typically applied directly to the outer support surface of the stent, possibly over all or part(s) of the inner surface of the stent, depending on how the coating system of the present invention described above is controlled, as shown in FIGS. 5A and 5B. The latter figure is a top and side area of filament 50 coated with coating material 52. Alternatively, the coating or coating mixture may be applied directly to the inner surface of the stent. A thin guide pin may be passed through one or more notches (i.e., windows) in the stent wall to allow the coating mixture to be applied directly to the desired location on the interior surface of the stent. In this way, it is possible to achieve coating of the filaments on the inside and outside with different coating materials containing different pharmaceutical ingredients. For example, the outer surface coating of the filaments may contain an anti-restenotic compound and the inner surface coating of the filaments may contain one of the second active agents mentioned above, such as an anti-thrombotic or anti-coagulant compound. If the stent has a sufficiently large diameter, thin "L-shaped" guide pins can be inserted into the open ends of the stent along the long axis of the stent to coat the inner surface of the stent.

Polymers useful in the present invention include, but are not limited to: poly (d, l-lactic acid), poly (d-lactic acid), ethylene vinyl alcohol copolymer (EVOH), epsilon-caprolactone, ethylvinyl hydroxylated acetate (EVA), polyvinyl alcohol (PVA), polyethylene oxide (PEO) and copolymers and mixtures thereof, dissolved in chloroform or acetone or other suitable solvents. All of these polymers have a safe and low history of inflammation for systemic circulation.

A non-polymeric coating may also be used in the present invention, such as everolimus coated onto the surface of a metal stent by ionic bonding.

Using the described coating system, it has been found possible to coat the surface of the stent, either on the top, on the side, or on the inside. By careful selection of the appropriate ratio of solvent to polymer, the viscosity of the solution can be adjusted so that some of the solution can move down the sides of the pillar and occupy the bottom surface before solidifying, see fig. 5B. By controlling the residence time of the catheter near the stent edge, the amount of polymer applied to the stent edge or base is increased or decreased. In the embodiment shown in fig. 3, a primer layer 34 composed of a neat polymer and a solvent is first applied to the stent surface 24 using the coating system of the present invention, and the solvent is then evaporated. A second polymeric coating 32 containing a bioactive agent is then applied.

As noted above, a second active agent may also be added to the polymer blend. Such as: the crystalline form of heparin may be added to the coating. Heparin crystals are micronized to a particle size of about 1-5 microns and then added to the polymer solution in the form of a suspension. Suitable forms of heparin when coated according to the process of the present invention are those crystalline forms that exhibit biological activity in mammals, including heparin salts (i.e., sodium heparin and low molecular weight heparins and salts thereof). When the drug-releasing stent is placed in the vessel wall, as shown in fig. 8, the heparin crystals near the surface of the coagulated polymer coating begin to dissolve, increasing the pores of the polymer. As the polymer slowly dissolves, more heparin and bioactive agent is released in a controlled manner.

However, referring to fig. 8, it should be noted that coating the inner surface of the stent is not always required. For example, coating the inner surface of the stent increases the corrugated conductive profile of the stent, making it less maneuverable within small vessels. Also, the inner surface is flushed directly by the blood flow through the stent after implantation in a blood vessel, resulting in loss of the drug released from the inner surface of the stent in the systemic circulation. Thus, in the embodiment shown in fig. 3 and 4, the coagulated polymer and active agent are mostly coated outside the circumference of the stent support, and secondarily on the sides. In a preferred embodiment, only a minimal amount of the polymer and active agent are coated on the inner surface of the stent. If desired, at least a portion of the interior surface of the stent may be uncoated or bare.

In addition, the coatings of fig. 3 and 4 may also be selectively applied to the surface of the stent wire. The thickness of the coating should correspond to the volume of bioactive coating used in the tissue. It is advantageous to limit the extent of the coating, since these extents of coating can lead to high damage when placing the stent.

A primer layer is first applied uniformly over the surface of the stent to promote adhesion of the bioactive agent-containing coating to the stent and/or to help stabilize the polymer coating on the stent. The initial coating may be applied by any method known in the art or by the precision feed system of the present invention. It is also within the scope of the invention to use different polymeric materials for the initial coating, such as parylene (poly (dichloro-p-xylene)), or any other material that will provide good adhesion between the metal substrate and the bioactive agent-containing coating. Parylene (poly (dichloro-p-xylene)) may be deposited using plasma deposition or vapor deposition techniques, which are well known in the art (see U.S. patent 6,299,604). In one embodiment of the invention, a discontinuous or continuous coating comprising heparin is applied to the inner surface of the stent and a coating comprising the aforementioned antiproliferative agents of the invention is applied to the outer surface of the stent.

When it is desired to form a coating on a metal stent substrate having a high drug/polymer substrate ratio, e.g., a drug content of 40-80% by weight of the coating, it is desirable to apply a primer layer to the stent filaments to stabilize and firmly adhere the coating to the substrate. The underlayer may be further treated by swelling in a suitable solvent such as acetone, chloroform, xylene or mixtures thereof prior to deposition of the coating material. This process is described in example 5 for the preparation of a stent with a high everolimus/poly-dl-lactide ratio.

Wherein a parylene base layer is formed on the stent filaments by a plasma deposition method, and then the base layer is swollen in xylene, and finally a coating material is deposited. The process can be used to produce coatings containing 50% drug or 75% drug in a poly-dl-lactide polymer substrate, with a coating thickness of only 5-10 microns.

As mentioned above, it is also within the scope of the present invention to produce a fully biodegradable stent using the coating system of the present invention. A tubular preform is first formed in the shape of the stent to be fabricated using an open-topped "C-shaped" helical channel into which polymer can be injected using a feed system. The outer diameter of the preform is open so that polymer can be placed into the preform, usually with one, if any more than one, passage of the injection tube; and at the same time establish a uniform edge of the stent, since the polymer will be constrained by the preform. The preform is soluble in a solvent, but the biodegradable scaffold produced thereby is insoluble in this solvent. After the polymer is injected and the solvent of the polymer solution has evaporated, the assembled assembly can be placed into a solution that dissolves the preform, releasing the intact scaffold structure. A representative starting material for making preforms is sucrose, which can be formed into the desired preform shape using standard injection molding techniques. A representative solvent for the preform is water.

Third section, methods of use, and performance characteristics

This section describes the vascular treatment method of the present invention and the performance characteristics of the stent made according to the present invention.

A. Method of producing a composite material

The methods of the present invention are directed to reducing the risk and/or extent of restenosis in patients at risk of local vascular injury or vascular occlusion. Typical vascular injuries are caused when a partially occluded vessel, such as a coronary artery or peripheral artery, is opened during an angiographic procedure. In angiography, a balloon catheter is placed over the occlusion and the balloon distal end is inflated and deflated one or more times, forcing the occluded vessel open. Such dilation of a vessel involves, inter alia, surface trauma to the vessel wall, which can dislodge plaque thereon, often causing local trauma sufficient to cause the vessel to respond with cell proliferation and reclosure over time. Not surprisingly, the incidence and severity of restenosis often correlates with the degree to which the vessels involved in angiography are pulled. Especially when overstretching to 35% or higher, restenosis occurs frequently and often severely, i.e. vessel occlusion.

In the practice of the present invention, the stent is placed at the distal end of the catheter in its contracted state, either within the lumen of the catheter, or at the distal end of the balloon in its contracted state. The distal end of the catheter is then guided to the lesion or potential occlusion point and the stent is released from the catheter, for example, by pulling a wire to release the stent to the site if the stent is self-expanding, or by balloon inflation on a balloon to expand the stent until the stent contacts the vessel wall tissue, thereby implanting the stent in the vessel wall tissue.

Fig. 6 shows a fully biodegradable stent and catheter for placing the stent into a vessel of the cardiovascular system, such as a coronary artery, according to the present invention. The stent 53 is shown in a partially relaxed state, referred to as a "drug helical stent". The stent is a self-expanding helical stent formed from polylactic acid and containing one or more bioactive agents of the present invention.

The helical stent is made from the preform, and the final expanded diameter of the preform is set to be slightly larger than the size of the lumen of the blood vessel to be treated with the helical stent. After removal of the preform, the delivery diameter of the helix was approximately 1/3 (at body temperature) of the final expanded diameter by twisting the ends of the drug helical stent in opposite directions to a smaller radius, thus compressing its full length under a slidable sheath. The thickness of the drug helical stent is small enough (about 25-125 microns) to allow it to bend easily to a smaller radius, forming a compressed helix comparable to the inner diameter of the sheath. The sheath is slidingly placed over a catheter 55 adapted to guide the stent in a compressed state into a target vessel. The sheath 54 has a handle 56 at its proximal end by which the angioplasty operator can pull the sheath back when the catheter tip reaches the proper location in the vessel to fully release the drug coil.

The catheter 55 has a central lumen of about 0.014 "diameter within which a guidewire 57 with a flexible tip 58 can slide. The catheter also has a luer hub 59 for connecting the lumen to the Y-site and hemostasis valve, as is well known in the angioplasty art. The outer diameter of the catheter with the slidable sheath ranges from 2-4F (french size); the peripheral artery would be larger if treated.

Since the drug helical stent is fully biodegradable, it does not interfere with the patient's subsequent opportunity to undergo uncomplicated vascular surgery, unlike an all-metal garment. In certain nerve and vascular applications, the placement of bare metal helical stents in blood vessels often results in thromboembolism, even total occlusion, and surprisingly, it has been discovered that the biocompatible polymer poly (dl-lactic acid) (PDLA) and mixtures thereof in the disclosed structure provide adequate mechanical strength to support the damaged vessel after angioplasty without causing embolism; and thus are exemplary materials for use in the production of the drug helical stents of the present invention.

Once in place, the scaffold begins to release the active compound into the cells of the vascular lining to inhibit cell proliferation. Fig. 7A shows the kinetics of everolimus release from two stents made according to the present invention, each having a coating (black squares) about 10 microns thick. The kinetics of drug release are achieved by immersing the stent in a 25% ethanol solution that greatly accelerates the rate of drug release from the stent coating. The graph shows the kinetics of drug release that can be expected in vivo, but over a much longer period of time.

Fig. 7B shows a graph of everolimus release from a coating of the present invention on a metal stent. The upper set of curves shows the drug release of the coating applied directly to the metal surface. The following set of curves (showing slow release) were obtained by first coating the surface of the metal stent with a base layer or primary coating of parylene followed by the coating system of the present invention. It can be seen that the primer layer increases the mechanical adhesion of the coating to the stent surface, slows the decomposition of the biodegradable coating, and also slows the drug release. When a strongly adherent stent coating is required to withstand the repeated abrasion of a drug eluting stent during bending operations within a catheter and/or vessel; and/or a slower rate of drug release at the implantation site following stent implantation is required to prolong the treatment of the atherosclerotic process.

Fig. 8 shows a cross-sectional view of a vascular region 60 into which a stent 62 having a coating 66 on coated wires such as wire 64 has been implanted. The figure shows the release of the anti-restenosis compound from each filament region into the surrounding vessel wall region. Over time, the smooth muscle cells of the vessel wall that are formed begin to grow within the stent or through the stent mesh or helical interstices, eventually forming a continuous inner cell layer engulfming both sides of the stent. If stent implantation is successful, the degree of occlusion of the vessel at the stent site will be less than 50% later, i.e., the diameter of the cross-section that maintains the blood flow passageway in the vessel is at least 50% of the diameter of the expanded stent when implanted.

Experiments conducted in animal models of porcine restenosis ("post balloon angioplasty restenosis- -a practical proliferative model of porcine coronary arteries", Circulation 82 (6) 2190-. The study is summarized in example 4.

In brief, several stents were compared in their extent of restenosis 28 days after implantation, including: bare metal stents, polymer coated stents containing high or low concentrations of rapamycin and everolimus.

Table 1 of example 4 shows that rapamycin (Rapa-high or Rapa-low) and everolimus (C-high or C-low) stents greatly reduced the level of restenosis with minimal restenosis for high dose everolimus stents. The same results were obtained in a low-damage animal study (table 2).

FIGS. 9A-9C are cross-sectional illustrations of stents for neointima 28 days after implantation of a bare metal S-stent (obtained from BiosensoresInternational Inc of Bewport Beach, Calif.). FIGS. 10A-10C are examples of neointimal formation of polymer coated (drug-free) S-stents. FIGS. 11A-11C and FIGS. 12A-12C are neointimal formations of everolimus/polymer coated stents. Overall, the vessels treated with everolimus-coated stents healed well, the endothelial layer was well established, and the vessels healed completely and reached vascular homeostasis for 28 days. FIG.13 is a cross-sectional view of a blood vessel at 91 times magnification showing the healing and formation of an endothelial layer within the lumen of the blood vessel 28 days after stent implantation.

The photographs show that the best binding to eliminate restenosis at 28 days is C-high or C-U high (see example 4), containing 325mg and 275mg everolimus, respectively, on a 18.7mm long stent. Data from follow-up visits to outbred piglets at 28 days predicted a 50% reduction in restenosis rate compared to bare metal stents (S-stents) currently on the market. The data also show that everolimus is superior to or at least equivalent to 180mg rapamycin on the same stent/polymer delivery platform. Morphometric analysis (example 4) supported these results.

Figure 15 shows the relationship between balloon overstretching of the vessel (measured by the balloon/to artery ratio, i.e. (B/a ratio)) and vessel damage in animal experiments. This data shows that creating highly controlled vascular injury with an over-expanded angioplasty balloon is a reasonably accurate method of creating predictable and known vascular injury in a pig model.

FIG. 14 shows a "best fit" linear regression curve of drug dose in polymer selected coated S-stents, with the degree of injury correlated to the area of stenosis at follow-up. The "stenosis area" is an accurate indicator of neointimal formation as determined by morphometric analysis. As can be seen from the figure, only a high amount of everolimus coated stents showed a negative slope with respect to the increase in the degree of damage in the test sample group. This analysis suggests that the C-high coating may be able to control restenosis in the damaged coronary artery, which is virtually independent of the degree of damage. None of the other coating formulations exhibit this unique characteristic.

From the foregoing, it will be seen that this invention is one well adapted to attain all the ends and objects set forth above, together with the advantages thereof.

In one aspect, the invention provides biodegradable stent coatings having a high drug/polymer ratio (e.g., 40-80% by weight drug). This property allows continuous release of the anti-restenosis compound from the low profile stent over a longer period of time. At the same time, the total amount of components of the polymer breakdown, such as lactate and lactic acid released during bioerosion, is relatively small, thereby minimizing possible side reactions, such as irritation caused by bioerosion of the stent coating.

In another aspect, the present invention provides an improved method of treating or inhibiting restenosis. The method involves a novel combination of macrocyclic triene immunosuppressive compounds in the polymer coating of a stent, which provides at least the same anti-restenosis effect as the best stents of the prior art, and additionally provides the added advantage over the prior art that the efficacy of the method appears independent of the extent of injury, and that the method can provide a greater degree of endothelialization of the stent vessels.

Finally, the present invention provides a fully biodegradable stent having the just mentioned advantages and having more flexibility in design than a metal stent, especially in terms of the total length of the stent and the future surgical possibilities of the treated vessel.

The following examples illustrate aspects of the fabrication and use of the stents of the present invention. The present invention is not limited to these ranges.

Example 1

Preparation of everolimus and derivatives thereof

Step A, Synthesis of 2- (tert-butyldimethylsilyl) oxyethanol (TBS ethylene glycol)

154ml of anhydrous THF and 1.88g of NaH were stirred under a nitrogen atmosphere in a 500ml round-bottom flask with condenser. 4.4ml of anhydrous ethylene glycol was added to the flask and stirred for 45 minutes to form a large amount of precipitate. 11.8g of tert-butyldimethylsilyl chloride were added to the flask and vigorous stirring was continued for 45 minutes. The resulting mixture was poured into 950ml of diethyl ether. The ether was washed with 420ml brine and the solution was dried over sodium sulfate. The product was concentrated by evaporation of ether in vacuo and then purified by flash chromatography on a 27X5.75cm silica gel column using hexane/Et₂O (75: 25v/v) solvent system. The product was stored at 0 ℃.

Step B, Synthesis of 2- (tert-butyldimethylsilyl) oxyethanol triflate (TBS ethylene glycol Trif)

Under nitrogen and vigorous stirring, 4.22g of TBS ethylene glycolAlcohol and 5.2g of 2, 6-lutidine were mixed in a 100ml two-necked flask equipped with a condenser. 10.74g of trifluoromethanesulfonic anhydride were slowly added over 35-45 minutes to give a yellowish brown solution. The reaction was then quenched by the addition of 1ml of brine and the solution was washed 5 times with 100ml of brine to a final pH of 6-7. The solution was dried over sodium sulfate and concentrated by evaporation of dichloromethane in vacuo. The product was purified by flash chromatography on a 24X3cm silica gel column with hexane/Et₂The O (85: 15v/v) solvent system was eluted and then stored at 0 ℃.

Step C, Synthesis of 40-O- [2- (tert-butyldimethylsilyl) oxy ] ethyl-rapamycin (TBS Rap)

In a 50ml flask, 400mg of rapamycin, 10ml of toluene and 1.9ml of 2, 6-lutidine were mixed and stirred at 55-57 ℃. In another 3ml septum vial was added 1ml of toluene followed by 940. mu.l of 2, 6-lutidine and then 2.47g of TBS ethylene glycol Trif. The mixture in the vial was added to a 50ml flask and the reaction was allowed to proceed for 1.5 hours with stirring. To the reaction flask was added 480. mu.L of 2, 6-lutidine and an additional 1.236g of TBS ethylene glycol Trif, and the reaction was continued with stirring for 1 hour. Finally, 480. mu.L of 2, 6-lutidine and 1.236g of TBS ethylene glycol Trif were added to the mixture, and the mixture was stirred for another 1-1.5 hours. The resulting brown solution was filtered through a vacuum fritted glass filter. The crystalline precipitate was washed with toluene until all color was removed. Then 60ml of saturated NaHCO was used₃The filtrate was washed 2 times with the solution and then with brine. The solution was dried over sodium sulfate and concentrated in vacuo. The product was dissolved in a small amount of hexane/EtOAc (40: 60v/v) solvent and purified by flash column chromatography on silica gel 33X2cm eluting with the same solvent. The solvent was removed in vacuo and the product was stored at 5 ℃.

Step D, 40-O- (2-hydroxy) ethyl-rapamycin (everolimus) synthesis process

A pyrex dish (150x75mm) was filled with ice and placed on a stir plate. A small amount of water was added. 60-65mg of TBS-rapamycin was first dissolved in 8ml of methanol in a small glass vial. Adding into a small bottle0.8ml of 1N HCl, the solution is stirred for 45 minutes and then 3ml of saturated NaHCO are added₃The aqueous solution is neutralized. To the solution was added 5ml brine, followed by 20ml EtOAc, resulting in a biphasic mixture. The two phases were mixed and the aqueous layer was drained off with a separatory funnel. The remaining solution was washed with brine until the final pH was 6-7 and then dried over sodium sulfate. The sodium sulfate was removed with a fritted glass filter and the solvent was then evaporated in vacuo. The resulting concentrate was dissolved in EtOAc/methanol (97: 3) and then purified by flash chromatography on silica gel 23X2cm eluting with the same solvent. The solvent was removed in vacuo and the product was stored at 5 ℃.

Example 2

Preparation of stents containing Everolimus in Poly-d, l-lactide coatings

100mg of poly-d, l-lactide were dissolved in 2ml of acetone at room temperature. 5mg everolimus was placed in a vial and 400. mu.L lactide solution was added. A syringe pump controlled by a microprocessor accurately feeds 10 μ L of the drug-containing lactide solution onto the top surface of the stent struts. The solvent evaporates to produce a uniform single drug-containing polymer layer on the stent.

In the same manner, 15 μ L of the solution was applied to the top and side surfaces of the stent struts, resulting in a single layer coating on the top and side surfaces of the stent struts.

Example 3

In vitro drug delivery from everolimus-containing stents in poly-d, l-lactide coatings

The coated stent was placed in 2ml phosphate buffer containing 25% ETOH, ph7.4, preserved with 0.05% (w/v) sodium azide and the temperature was maintained at 37 ℃ to construct in vitro drug release. All buffers were periodically withdrawn for sampling for drug measurements, while the same volume of fresh buffer (infinite pellet) was added. Fig. 7 illustrates drug release from 2 similar stents coated with a single polymer coating in this manner.

Example 4

Animal implantation experiment

A. QCA results of porcine safety and dose studies

The basic principle is as follows:

as is known, the most challenging condition for drug eluting stent treatment is severely damaged vessels, since it is known that an increase in the extent of vessel damage directly leads to an increase in the extent of restenosis (neointima formation). The experiment was performed in pigs and the target vessel of the drug-coated stent was severely damaged with an angioplasty balloon (mean vessel overstretch damage of about 36%). This causes severe tearing and stretching of the intimal and medial layers of the vessel, resulting in extravasation restenosis after implantation of the stent 28. In this way, the relative effectiveness of various doses of drug and different weight ratios of drug to polymer in reducing restenosis at 28 days post-stent implantation can be evaluated on the same metal stent/polymer platform.

Abbreviations：

"bare stent" refers to an 18.7mm bare metal corrugated ring stent (i.e., the "S-stent" currently marketed by biosensorsintl.inc.);

"C-high" refers to an 18.7mm stent with 325mg of everolimus in the PDLA (poly-dl-lactic acid) polymer coating.

"C-low" refers to an 18.7mm stent with 180mg everolimus in the PDLA polymer coating.

"rapamycin-high" refers to an 18.7mm stent containing 325mg rapamycin in the PDLA polymer coating.

"rapamycin-low" refers to an 18.7mm stent containing 180mg of rapamycin in the PDLA polymer coating.

"C-U high" refers to an 18.7mm stent containing 275mg everolimus in a very thin PDLA polymer coating (37% drug to polymer by weight).

"C-U Low" refers to an 18.7mm stent containing 180mg everolimus or equivalent (37% drug to polymer weight) in a very thin PDLA polymer coating.

"Polymer scaffold" refers to an 18.7mm S-scaffold with only PDLA polymer coating.

"B/A": the ratio of the final expanded balloon to the arterial vessel indicates the degree to which the vessel is overstretched.

"Mean Luminal Loss (MLL)": the stent lumen was measured 3 times during implantation, and the difference between the average of the 3 data minus the average of the 3 measurements during follow-up imaging reflected the amount of neointima formation in the stent.

Method of producing a composite material：

A drug-eluting stent consisting of a corrugated-ring wire mesh stent (i.e., S-stent) and a polymer coating was implanted into distant breeding piglets (also jekata piglets could be implanted for more than 28 days), with different doses of everolimus or rapamycin applied to the coating. At the time of implantation, the blood vessel diameters before and after implantation of the stent were measured by QCA (quantitative coronary angiography). Animals were re-imaged in the mount area before euthanasia at 28 days or longer as specified in the table below.

After the animals were euthanized according to the approved protocol, the hearts of the animals were removed and injected under pressure into the coronary arteries with formalin. The coronary artery portion containing the stent was removed from the surface of the heart and then fixed to an acrylic plastic plate and the cross-section was cut with a diamond saw. Sections of 50 micron thick acrylic material containing the most proximal, middle, and most distal vessels were optically polished and mounted on slides of a microscope.

A microscope with a digital camera was used to take high resolution images of the cross-section of the blood vessel mounted on the slide. The images were analyzed for tissue morphology according to the following procedure.

The analysis of the morphology of the tissues was performed by a PC system based a.g. heinze microscope using the computer Image processing system Image Pro Plus 4.0.

1. Average cross-sectional area and lumen thickness (area delineated by intima/neointima-lumen edge); neointima (the area between the lumen and the inner elastic lamina, IEL, when the IEL disappears, the area between the lumen and the residual media or outer elastic lamina, EEL); middle (area between IEL and EEL); vessel size (area bounded by EEL, but not adventitial area); adventitial area (the area between the peripheral adventitial tissue, adipose tissue and myocardium and the EEL).

2. Degree of damage. To quantify the extent of vascular injury, a score based on the length and amount of different vessel wall structure tears was used. The damage degree is calculated as follows:

0 ═ complete IEL;

1 ═ IEL was slightly torn, exposing the epilayer (minor injury);

2-IEL moderately torn, exposing the deeper middle layer (moderate cut);

the 3-EEL was torn, exposing the outer layer. mm is²

The following table is the results of the QCA analysis (measurement of average lumen loss due to restenosis). The data in the "neointimal area" column below shows the results of a morphological analysis of blood vessels and stents removed from pigs at follow-up

TABLE 1 results of "high Damage" experiment

Description of the apparatus	(B/A) ratio (average)	Time of follow-up	Average inner cavity loss (mm)	Neointima area (mm)2)	Bracket number
						Bare metal stent	1.33	28	1.69	5.89	31，39，40，45，47，50
Polymer coated stent	1.36	28	2.10	5.82	32，41，43，48，51，60
						R-is high	1.39	28	1.07	3.75	42，44，49，65，69，73
R-low	1.42	28	0.99	2.80	52，56，61，64，68，72
						C-high	1.37	28	0.84	3.54	54，55，59，63
C-low	1.36	28	1.54	3.41	53，57，58，62，66，70，74
						High C-U	1.36	28	0.85	2.97	67，75，92，103

B. Low damage study:

to further determine the optimal dose of everolimus in mildly damaged vessels, particularly in patients with simple coronary artery disease and single untreated lesions, implantation of everolimus drug eluting stents resulted in moderate to low overstretch injury (about 15%). Experiments were performed on farm pigs for 30 days and safety studies were performed on adult eucatata piglets for 3 months of implantation. The angiographic results were as follows:

TABLE 2 QCA results of the "Low Damage" experiment

Description of the apparatus	(B/A) ratio	Days after implantation	Average inner cavity loss (mm)	Neointima area (mm)2)	Bracket number
						Bare metal stent	1.14	28	0.95	2.89	20，22，26，29
Bare metal stent	1.13	90			76，80，84，87，91
						High C-U	1.15	28	0.60	2.14	94，96，98，102
C-U is low	1.09	28	0.49	2.26	93，95，97，100，101
						High C-U	1.15	90			77，81，85，86，90

The above data suggest that low or high doses of C-U of everolimus can reduce neointimal formation by 45-48% in low to moderate damaged vessels.

C. Morphometric analysis

The percent stenosis area was calculated by computer measurements of the total cross-sectional area within each stent and the cross-sectional area of new tissue (neointima) formed within the stent. The average degree of vascular damage, neointimal area and percent stenosis area (three sheets per stent measured, averaged) for each drug and polymer formulation is shown in the table below.

TABLE 3 "high Damage" test results

Description of the apparatus	Degree of damage	Time of follow-up	Neointima area (mm)2)	Narrow area (%)	Bracket number
						Bare metal stent	1.9	28	5.89	0.72	31，39，40，45，47，50
Polymer coated stent	2.11	28	5.82	0.70	32，41，43，48，51，60
						R-is high	2.10	28	3.75	0.55	42，44，49，65，69，73
R-low	1.90	28	2.80	0.43	52，56，61，64，68，72
						C-high	1.89	28	3.54	0.38	54，55，59，63
C-low	2.1	28	3.41	0.53	53，57，58，62，66，70，74
						High C-U	2.13	28	2.97	0.45	67，75，92，103

Morphometric analysis is a highly accurate method of measuring in-stent restenosis in porcine coronary models. In the high injury model, the C-high formula produced the lowest amount of neointimal formation at 28 days of the high injury experiment; while the C-U high group had the highest degree of injury, but still maintained a low percentage of narrow area-0.45. Thus, this data independently confirms that the findings from the QCA analysis, the preferred formulation should be C-U high to support human experiments.

D. Histological analysis

Slides of C-U high and rapamycin-low were presented to an experienced cardiologist reviewing inflammation, fibrin, endothelialization of newly healing vascular lumens in the cross section of the blood vessels. No difference in histological changes caused by rapamycin and everolimus eluting stents was found. Overall, the blood vessels had formed an intact endothelial layer at 28 days to restore their equilibrium and complete healing was achieved. FIG.13 is an example of a cross-sectional view of a blood vessel at 91 times magnification showing healing of the vessel lumen and establishment of the endothelial layer 28 days after stent implantation.

E. Comparison with published results

Carter et al have published the results of experiments on porcine rapamycin coated stents using a Palmaz Schatz metal stent. The following table compares the published results for Carter and the results from the Biosensors experiments: mm is

TABLE-4

Description of the apparatus	Blood vessel overstretching (%)	Average late loss (mm)	Standard deviation (mm)	Section of neointimaArea of surface (mm)2)
					S-Stent bare Metal control	33.5％±9.2％	1.80	±0.5	7.6
S-Stent coating with Polymer only	34.9％±4.8％	2.02	±0.8	8.5
					S-Stent Polymer/rapamycin (325mg)	32.9％±10.1％	0.66	±0.2	3.27 (control-57%)
S-scaffold Polymer/Everolimus (325mg)	36.8％±8.5％	0.74	±0.3	3.61 (control-50%)
					PS scaffold BARE＊Control	10-20％	1.19	---	4.5
PS Stent OnlyCoating polymers	10-20％	1.38	---	5.0
					PS rapamycin eluting stent＊166mg	10-20％	0.70	---	2.9 (control-33.5%)
PS rapamycin eluting stent＊166mg (Slow release)	10-20％	0.67	---	2.8 (control-37.7%)
					PS rapamycin eluting stent＊450mg	10-20％	0.75	---	3.1 (control-31.1%)

Example 5

Preparation of high drug content stents

A commercially available 14.6mm long metal undulating ring stent (Biosensors Intl S-stent, undulating ring design) was first coated with a approximately 2 micron thick bottom layer of parylene "C" by plasma deposition. The parylene coated stent was then placed in xylene at ambient temperature overnight. A poly-d, L-lactic acid stock solution containing 50. mu.g/. mu.L of PDLA was prepared by dissolving 100mg of polylactic acid (PDLA) in 2ml of acetone.

To prepare a coated stent with a drug to polymer ratio of 50%, 5mg everolimus was dissolved in 100 μ L PDLA stock solution. An additional 20. mu.L of acetone was added to facilitate dispersion of the solution. The scaffolds were removed from the xylene and the solvent carefully dipped dry. A total of 5.1 μ L of coating solution was dosed to the outer surface of each stent. The scaffolds were dried at room temperature and placed in a desiccator to dry overnight. Thus, each scaffold contained 212. mu.g of everolimus in 212. mu.g of PDLA.

To prepare a coated stent with a drug to polymer ratio of 75%, 5mg everolimus was mixed with 33.3 μ L PDLA stock solution. An additional 33.3. mu.L of acetone was added to dissolve the mixture. The scaffolds were removed from the xylene and dip-dried as described above. A total of 2.8 μ L of coating solution was dosed to the outer surface of each stent. The scaffolds were dried at room temperature and placed in a desiccator to dry overnight. This resulted in 212. mu.g everolimus per stent in 70. mu.g PDLA.

The everolimus/PDLA coating thickness of the finished stent was about 5 microns or exhibited a light milky appearance with smooth top and side surfaces and strong connections to the metal strut surfaces.

Claims (18)

1. An intravascular stent for placement at a site of vascular injury to inhibit restenosis at the site, comprising:

a radially expandable tubular body formed of connected filaments, each filament having a top surface, side surfaces and an interior support surface, an

A biodegradable drug-releasing layer comprising a restenosis-inhibiting macrocyclic triene immunosuppressive drug and a biodegradable polymer, wherein the drug is present in an amount of 40 to 80% by weight of said drug-releasing layer, said drug-releasing layer being coated only on the top surface and at least a portion of the side surfaces of said filaments but not on the inner surface thereof.

2. The stent of claim 1, further comprising a bottom layer positioned between the filaments and the drug release layer.

3. The stent of claim 2 wherein the drug-releasing layer is comprised of (i) 20-60% by weight of a poly-d 1-lactide polymer substrate and (ii) 40-80% by weight of a restenosis-inhibiting compound, and the base layer is a polymeric base layer.

4. The stent of claim 1, wherein the filaments are made of metal.

5. The stent of claim 1, wherein the compound has the structure

Wherein R is H or CH₂-X-OH and X is a linear or branched alkyl group containing from 1 to 7 carbon atoms.

6. The stent of claim 5, wherein X is-CH₂。

7. The stent of claim 1 wherein the inner surface of the stent body is coated with a second drug-releasing layer comprising a second drug.

8. The stent of claim 1, wherein the depth of the drug release layer on the top surface of the filaments is non-uniform.

9. The stent of claim 1, wherein the tubular body is biodegradable.

10. The stent of claim 1, wherein the tubular body is made of poly-L-lactic acid.

11. A biodegradable intravascular stent for placement at a site of vascular injury to inhibit restenosis at the site, comprising:

a radially expandable tubular body formed of connected filaments, the tubular body being made of poly-L-lactic acid, and

a biodegradable drug-releasing layer comprising a restenosis-inhibiting macrocyclic triene immunosuppressive drug and a biodegradable lactic acid copolymer, wherein the drug content is 40 to 80% by weight of said drug-releasing layer, and the polymer content is 20 to 60% by weight of said drug-releasing layer.

12. The stent of claim 11, further comprising a bottom layer positioned between the filaments and the drug release layer.

13. The stent of claim 12 wherein the drug release layer is comprised of (i) 20-60% by weight of poly-dl-lactide polymer substrate and (ii) 40-80% by weight of a restenosis inhibiting compound, and the base layer is a polymeric base layer.

14. The stent of claim 11, wherein the drug release layer has a thickness of 3 to 20 microns.

15. The stent of claim 11, wherein the compound has the structure

Wherein R is H or CH₂-X-OH, and X is linear or branchedAlkyl groups having 1 to 7 carbon atoms.

16. The stent of claim 15, wherein X is-CH₂。

17. The stent of claim 11 wherein the inner surface of the stent body is coated with a second drug-releasing layer comprising a second drug.

18. The stent of claim 11, wherein the depth of the drug release layer on the top surface of the filaments is non-uniform.