US20040024448A1 - Thermoplastic fluoropolymer-coated medical devices - Google Patents

Thermoplastic fluoropolymer-coated medical devices Download PDF

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Publication number
US20040024448A1
US20040024448A1 US10/213,126 US21312602A US2004024448A1 US 20040024448 A1 US20040024448 A1 US 20040024448A1 US 21312602 A US21312602 A US 21312602A US 2004024448 A1 US2004024448 A1 US 2004024448A1
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United States
Prior art keywords
coating
implantable device
stent
copolymer
therapeutic agent
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Abandoned
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US10/213,126
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English (en)
Inventor
James Chang
Robert Cleek
Edward Cully
Michael Vonesh
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WL Gore and Associates Inc
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Gore Enterprise Holdings Inc
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Publication date
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Priority to US10/213,126 priority Critical patent/US20040024448A1/en
Assigned to GORE ENTERPRISE HOLDINGS, INC. reassignment GORE ENTERPRISE HOLDINGS, INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: CHANG, JAMES W., CLEEK, ROBERT L., VONESH, MICHAEL J., CULLY, EDWARD H.
Priority to EP03767251A priority patent/EP1545642B1/en
Priority to AU2003258112A priority patent/AU2003258112B2/en
Priority to CA002493634A priority patent/CA2493634C/en
Priority to AT03767251T priority patent/ATE423579T1/de
Priority to DE60326367T priority patent/DE60326367D1/de
Priority to PCT/US2003/024618 priority patent/WO2004012783A1/en
Priority to JP2004526062A priority patent/JP5048216B2/ja
Publication of US20040024448A1 publication Critical patent/US20040024448A1/en
Priority to US11/378,401 priority patent/US8048440B2/en
Priority to US13/282,892 priority patent/US8609125B2/en
Assigned to W. L. GORE & ASSOCIATES, INC. reassignment W. L. GORE & ASSOCIATES, INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: GORE ENTERPRISE HOLDINGS, INC.
Priority to US14/107,343 priority patent/US20140105955A1/en
Abandoned legal-status Critical Current

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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
    • A61L31/08—Materials for coatings
    • A61L31/10—Macromolecular materials
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
    • A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
    • A61L27/28—Materials for coating prostheses
    • A61L27/34—Macromolecular materials

Definitions

  • the present invention relates to the field of medical devices provided with at least a partial surface coating of polymer.
  • Surgical interventions often involve the implantation of a medical device, typically manufactured from polymeric and/or metallic materials, that is intended to provide a mechanical repair of a medical malady. While providing necessary and often life saving benefits, the implanted metal or polymer material may also produce some type of complication. Some of the more common complications include acute thrombosis; increased risk of infection immediately post procedure and/or chronically; fibrous encapsulation of the device resulting from a foreign body response and inflammation; and vascular proliferative disease resulting in an excessive, inflammatory, fibroproliferative response to injury.
  • therapeutic agents are administered to ameliorate complications arising from the medical implant and the disease being treated. Most often these are administered orally or through injection and result in systemic delivery. Ideally therapeutic agents would be released locally in a controlled fashion from an implant to maximize the effectiveness of the agent at the desired site without causing severe systemic side effects.
  • a combination device, or product, that provides for local drug delivery and a mechanical solution to the medical malady may result in clinical outcomes not possible otherwise.
  • One approach to achieving this combination is through the use of coatings applied to the surfaces of medical devices, implantable for short or long terms, wherein the coating may optionally contain therapeutic agents elutable from the coating.
  • biocompatible materials that can adequately retain an efficacious dose, provide for prolonged drug release, and be incorporated into the mechanical device, in the simplest possible fashion, without compromising the device functionality.
  • the material would truly be exemplary if it provided more benefits to the combination product than functioning solely as a matrix for the release of a therapeutic agent.
  • this can be accomplished without the addition of still another component, such as an adhesive material or primer coatings, or without requiring surface modification of the medical device, but rather with the polymer material itself serving as a biocompatible adhesive with or without additives.
  • biodegradable materials for drug delivery such as alpha hydroxy esters
  • these compounds have glassy or rigid amorphous states that do not meet the flexibility requirements of combination implantable device.
  • These materials have poor adhesive properties, particularly with regard to common materials used to manufacture medical devices such as various metals and polymers such as polytetrafluoroethylene (PTFE).
  • PTFE polytetrafluoroethylene
  • Silicones are among the most widely used synthetic polymers that are intended to be non-biodegradable and are found in a variety of medical applications. They are sometimes used as a matrix material for elution of therapeutic agents, and as an elastomer they offer a good degree of flexibility. See, for example, U.S. Pat. No. 6,358,556 to Ding et al. Silicones consist of at least three components: an elastomer, silica reinforcing agent, and a volatile inhibitor to stop cross-linking. However, silicones have poor bonding strengths to many medical device substrate materials, and poor long-term in vivo tensile strength. They are less biocompatible than most fluoropolymers.
  • Silicones absorb lipids and proteins over time, have a tendency to generate particulate debris over time, and exhibit poor abrasion resistance. Curatives in the vulcanized polymer can be problematic in that they may react with additive. Other problems are known to include cracking, swelling (generally due to lipid or protein absorption), tear propagation and poor adhesion. These problems are exacerbated by the use of additives.
  • Various fluoropolymer materials have been proposed as drug delivery material; see, for example, EP 950386 to Llanos et al. which suggests a list of materials including PTFE. While PTFE is particularly inert and highly biocompatible, it is not elastomeric and is limited in elution capability if not used in its porous expanded form (ePTFE). Drugs are typically eluted from the interconnected void spaces of ePTFE rather than by molecular diffusion from within the polymer matrix; see, for example, U.S. Pat. No. 5,290,271 to Jernberg.
  • EP 1192957 to Llanos et al. proposes other fluoropolymer materials comprising a first monomer chosen from the group consisting of vinylidene fluoride and a second monomer that is different from the first monomer.
  • These materials are relatively non-durable according to examples that describe cracking of the matrix during device expansion.
  • these particular materials are limited in their drug loading and drug elution capabilities. The ability to bond to a variety of other materials without requisite primer coating or surface treatment of the substrate, the ability to function as an integral component of a coated medical device (without adverse effect on the device function), and the ability to aid in the manufacturing of a wide range of combination products has not been shown
  • the present invention provides a coating for a medical device that can also act as a vehicle for delivery of therapeutic agents.
  • the coating comprises a fluoropolymer that is highly inert and biocompatible, has elastomeric characteristics that provide desirable mechanical properties such as good flexibility and durability, can be loaded with additives (such as therapeutic agents) either in solid or liquid form, and as such has excellent drug elution characteristics.
  • the coating material is a thermoplastic copolymer of tetrafluoroethylene (TFE) and perfluoroalkylvinylether (PAVE) that is free of cross linking monomers and curing agents.
  • the coating material is preferably an amorphous thermoplastic.
  • the perfluoroakylvinylether may be perfluoromethylvinylether (PMVE), perfluoroethylvinylether (PEVE) or perfluoropropylvinylether (PPVE).
  • PMVE perfluoromethylvinylether
  • PEVE perfluoroethylvinylether
  • PPVE perfluoropropylvinylether
  • the coating material is considered “functionally transparent” for most medical device coating applications. This means that the coating accomplishes its intended purposes without adverse effect on the intended function of the coated medical device.
  • the coating material provides this unique feature in its ability to withstand mechanical deformations required for the assembly, deployment, expansion, and placement of medical devices, to such an extent that the user of such devices does not realize that the material is present.
  • the coating is functionally transparent to the coated medical device even while containing additives and providing for controlled release of therapeutic additives, if desired.
  • the coating material has good adhesive properties, such that it does not require a primer or other surface treatment of the substrate to be coated, and thus, can allow for thin coatings with or without additives that can also be used to bond components of the device together.
  • a medical device is hereby defined as any device used in the cure, mitigation, treatment, or prevention of disease, in man or other animals, or intended to affect the structure or any function of the body of man or other animals.
  • Medical devices are typically used in contact with any body fluids or body tissues of man or other animals.
  • Implantable medical devices are those devices that are inserted into living bodies for appreciable periods. More specifically, long term implants are defined as items implanted for more than 30 days while devices inserted into living bodies for lesser periods are considered to be short term implantable devices.
  • additives are considered herein to be any additional materials added to the TFE/PAVE copolymer for any reason, regardless of form. They may therefore be in the form of liquids or solids; they may represent solutions (including colloidal suspensions), mixtures, blends, particulates, etc.
  • amorphous is used herein to define a polymer that is substantially non-crystalline, and in which the molecular chains exist in the random coil conformation, with little or no regularity of structure.
  • the copolymer has sufficient amounts of PAVE in the molecular chains to substantially disrupt the crystallinity of the resulting fluoropolymer. Crystallinity can be detected by thermal/calorimetric techniques which measure the latent heat of the melting/freezing transitions. One convenient method of detection known to those of skill in the art is by Differential Scanning Calorimetry (DSC).
  • DSC Differential Scanning Calorimetry
  • the heat of fusion calculated from any endotherm detected in a DSC scan for the as-polymerized copolymer is no more than about 3 J/g, and preferably no more than about 1 J/g.
  • the scan rate should be set at 10° C. per minute over a temperature range beginning at 60° C. and ending at 400° C.
  • thermoplastic is used herein to define a polymer that will repeatedly soften when exposed to heat and return to its original condition when cooled to room temperature. Such a polymer can be made to soften, flow or take on new shapes, without significant degradation or alteration of the polymer's original condition, by the application of heat or heat and pressure.
  • a thermoplastic is accordingly a polymer in which the molecular chains are held together by the secondary van der Waals bonds; when enough thermal energy is applied, the chains break free from one another and the material will flow and melt.
  • thermoset polymer In contrast to a thermoplastic polymer, a “thermoset” polymer is hereby defined as a polymer that solidifies or “sets” irreversibly when cured.
  • Thermoset polymers have a three-dimensional network structure which prevents chains from being freed at higher temperatures. They will typically burn instead of melt.
  • a determination of whether a polymer is a “thermoplastic” polymer within the meaning of the present invention can be made by slowly elevating the temperature of a slightly tensioned specimen and watching for deformation. If the polymer can be made to repeatedly soften, flow, or take on a new shape, without significant degradation or alteration of the polymer's original chemical condition, then the polymer is considered to be a thermoplastic. If only small amounts of material are available it may be necessary to use a hot stage microscope for this determination.
  • Stents including both balloon expandable and self-expanding stents, are particularly improved by coating with the fluoropolymer.
  • the change of overall dimensions of the stent is accommodated by the good flexibility and durability of the coating material.
  • Coated stents of the present invention can be used for applications in vascular and non-vascular body conduits such as biliary, hepatic or esophageal.
  • the flexibility and adhesion of the coating substantially reduces or eliminates risk of cracking during stent expansion.
  • the durability of the coating reduces risk of damage to the coating by a stent delivery catheter or by the luminal surface of the body conduit into which it is inserted. This is of particular utility during single balloon procedures when the balloon is inserted into a lesion together with a stent, wherein the balloon simultaneously expands the stent and forcibly opens up the stenosis at the device deployment site.
  • Stent-grafts that is, stents that are provided with a covering, often of a tubular graft material, that covers some or all of the otherwise open interstices of a deployed stent, can also be provided with a coating over the surfaces of the stent and/or the surfaces of the graft material with beneficial results.
  • the tubular graft material is most typically PTFE, PET or polyurethane.
  • the coating may be used as an adhesive to join the graft covering to the stent. Likewise, it may be desirable to cover surfaces selectively so that only some surfaces are covered, or some surfaces are only partially covered. Further, the coating material itself may be used in the form of a thin film as the graft covering material.
  • the film may be applied in the form of thin sheets, tapes or tubes to the desired surface of a stent to create a stent-graft, covering the stent elements and the interstices between adjacent elements.
  • a stent may be dipped into the coating to achieve a covering that covers the stent elements and spans the stent interstices, thereby resulting in a stent-graft
  • the coating may be used with or without additives.
  • the coating may be used beneficially without an additive by covering a less biocompatible material, in effect passivating the less biocompatible material.
  • both stent elements and graft coverings for stent-grafts may be beneficially covered by the inventive coating (with or without additives) without significant adverse affect on device profile (for thin coatings) and without adverse effects on the coating such as cracking during device deployment, the coating is deemed particularly useful for all bendable elements of medical devices including both stent elements and graft coverings for stent-grafts. Bendable elements are considered to be those elements of a medical device that undergo bending during insertion into or use with a body.
  • Expandable endoluminal elements are considered to be those portions of an expandable device such as a stent that undergo a change of dimension during the course of the expansion of the device, from its initial shape appropriate for insertion into a body and transfer to a desired site, to its deployed size at the desired site.
  • Devices such as vascular grafts, venous valves, heart valves and heart valve leaflets, left ventricular assist devices, ocular implants including lenses and corneal implants, device introducers, access ports, topically-applied devices (e.g., wound dressings and transdermal patches), embolic filters, embolic particles, catheters, device delivery components, catheter balloons, guidewires, occluders, implantable electrical leads and devices, implantable patches including vascular and hernia patches, sutures and other surgical fasteners, and orthopedic implants can be beneficially coated.
  • Catheter balloons for stent delivery can be beneficially coated to improve their ability to retain a stent during insertion of a stent delivery system into a body conduit, substantially reducing any risk of loss or misplacement of a stent during the insertion process due to the stent having inadvertently moved with respect to the balloon surface. This risk can be considerable when it is attempted to insert a catheter balloon and stent into a restrictive vascular lesion.
  • Vascular grafts can be provided as tubular grafts or as sheet grafts for the repair of only a portion of the circumference of a blood vessel. This list is intended only to be representative of the types of medical devices that may be improved by the present invention, and consequently is not limiting. Further, the coating (with or without an additive) may be used as an adhesive between different components of a medical device.
  • the medical devices can be coated by a variety of known processes including spraying, dip-coating, powder coating, dispersion coating, lamination to other substrates, extrusion, molding, compression molding, or any other suitable means. It can be applied as very thin coatings, even when loaded with additive materials, and as such it enables medical devices to be made with minimal effect on the thickness and profile of the devices.
  • the coating material adheres well to a variety of substrates including various metals (e.g., stainless steels and nitinols) and to various polymers (e.g., ePTFE). It does not require special preparation of the substrate surface, additional bonding agents, or high temperature processing.
  • the coating material adheres well and because it has an elastomeric character, it is effective for use on stents and stent-grafts.
  • the coating is not adversely affected by the stent deployment process involving expansion of the stent diameter from its small, compacted diameter at which it was inserted into the vasculature, up to its larger diameter following deployment and expansion.
  • the coating does not crack or otherwise disrupt during this expansion process, which may involve plastic deformation of the metallic stent elements.
  • Any elutable therapeutic agent contained in the coating can thus be expected to follow its intended release rate because of the robust and durable character of the coating, minimizing risk of cracking or loss of adhesion.
  • the coating may be employed in the manufacture of medical devices as a drug-eluting adhesive.
  • the coating material is self-adhesive, meaning that additional layers of the coating will adhere well to previously applied layers.
  • the inventive coating material can be a practical, low temperature adhesive. It is generally preferable to use solvents of this type that typically do not dissolve or chemically react with most additives. A uniform coating of a heterogeneous mixture of a drug and the copolymer is possible with the present invention. This allows for coexistence of drug-loaded regions adjacent to drug-free regions in a pattern design that most efficiently delivers a drug (or other therapeutic agent) in a localized, strategic fashion. Thus, one common solvent and polymer are used to easily create polymer-drug regions in contact with polymer-coated regions that do not contain drugs.
  • solvents such as FC-75 fluorinated solvent (3M Fluorinert, 3M Specialty Chemicals Division, St. Paul, Minn.
  • the coating may be applied over any or all surfaces of a medical device.
  • the coating can be provided over the entire surface area of a medical device in a fully continuous fashion whereby none of the original surface of the device remains exposed. Alternatively, only some surfaces may be covered or some surfaces may only be partially covered.
  • the coating may be provided in discontinuous fashion such that it is interrupted at desired portions of a surface, for example, the coating may be provided as a dot-matrix pattern on a desired surface.
  • the coating can be provided as a film, in the form of thin sheets or tubes, in which form it can be used for numerous applications.
  • the film can be used as the covering material over various devices (including, as previously mentioned, a stent to create a stent-graft). It may be used as a stand-alone biological barrier material, for example, to separate different types of living tissues during healing. It can be applied over other substrates and subsequently bonded to the substrate by the use of heat or by the use of more of the coating polymer in a liquid form as an adhesive.
  • Films can be made with conventional methods including extrusion and solvent casting. They can be separately made for subsequent application to the surface of a device, or alternatively can be provided by techniques such as dip-coating directly over the surfaces of various devices (such as, for example, guidewires and stents).
  • the coating can be provided over porous substrates in order to reduce their porosity and/or permeability, including to an extent that the porous substrate is rendered non-porous across opposing surfaces of a device made from the porous substrate.
  • the coating can be provided in porous forms.
  • the copolymer material may be rendered porous by methods such as the inclusion of foaming agents, dissolving impregnated particles or by forcing gasses or supercritical fluids through the thermoplastic coating.
  • the coating material can be provided with a wide variety of additives including a variety of therapeutic agents.
  • the additive can remain stable and resident with the coating material (e.g., radiopaque additives), or alternatively an additive can be provided to allow for its elution over a specific period.
  • Solid or liquid additives may be used with the coating material at the time of coating the medical device surface. While smaller particle sizes are preferred for particulate additives, the coating material can accommodate larger particle sizes with minimal effect on the mechanical properties of the coating.
  • Additives may be provided in relatively high weight percent amounts, such as about 1, 2, 5, 10, 20, 30, 40, 50, 60, or 70 weight percent, or more.
  • Elutable additives may be provided with an additional capping layer of the coating polymer in order to reduce the rate of elution and extend the time of elution.
  • the capping layer may be provided over another filled layer of the coating, or alternatively may be applied directly over a layer of the additive material itself in some instances.
  • a capping layer may contain an additive which is different from an additive contained beneath the capping layer.
  • the release kinetics may be varied in other ways, such as, for example for particulate additives, by controlling the particle size as well as the weight percent loading.
  • additives may be used that are thermally activated, or that enhance in vivo imaging during, for example, fluoroscopic or magnetic resonance imaging. These latter additives are referred to herein as imaging opaque substances. Radioactive additives may be used to locally deliver radiation therapy.
  • bioactive therapeutic agents such as antithrombotic drugs including heparin, paclitaxol, dexamethasone and rapamycin are most commonly proposed to aid the performance of stents, many others can also be used beneficially, either alone or in various combinations.
  • Therapeutic agents for a wide variety of applications can be used as additives with the coating for use with various devices. These agents include, but are not limited to, antithrombotic agents, anticoagulants, antiplatelet agents, thrombolytics, antiproliferatives, antiinflammatories, hyperplasia and restenosis inhibitors, smooth muscle cell inhibitors, antibiotics, antimicrobials, analgesics, anesthetics, growth factors, growth factor inhibitors, cell adhesion inhibitors, cell adhesion promoters and drugs that may enhance neointimal formation such as the growth of endothelial cells.
  • these therapeutic agents may be used alone or in various combinations, and may be in coatings that cover all surfaces of a device or only portions of a device.
  • Additives that are not bioactive and not elutable can be used, for example, various pigments, MRI-opaque additives or radiopaque additives (e.g., tantalum or gold) that are used to enhance imaging of an implanted device. Encapsulated void spaces may be used for enhanced echogenicity during procedures such as ultrasound. Pigments may be beneficially added to enhance direct visualization, for example, to provide a contrast against the blood of a surgical field. Pigments may also be used for printed indicia for various labeling or instructional purposes. Specialty pigments (e.g., luminescent) may be used for particular applications, such as enhancing visibility of devices (e.g., guidewires) in darkened catheter labs.
  • devices e.g., guidewires
  • pockets or layers of an additive may be captured within the coating. These pockets or layers may then be exposed to body fluids by penetrating the coating with a sharp surgical device or tool.
  • elution as described herein pertains to diffusion of an additive that occurs within a solvent, where the solvent may be any suitable fluid including body fluids.
  • the solvent may be any suitable fluid including body fluids.
  • the copolymer When the additive is contained within a copolymer, such as for controlled release within the body, the copolymer must wet in order for elution to take place.
  • diffusion is defined to mean the transport of matter by random molecular motion from one region in space to another. It is one of the processes that govern the elution of additives from a copolymer-additive formulation.
  • the coating material of the present invention can be exposed to high temperatures without degradation.
  • the coating material can be exposed to 330° C. for one hour with a resultant weight loss of less than five percent, preferably less than one percent, and more preferably less than 0.5 percent. Because of the high temperature capability and the inert character of the coating material, it lends itself to high temperature sterilization. Because it is not hydrolyzable, it is not adversely affected by conventional steam sterilization. Further, sterilization does not adversely affect elution profiles as long as the sterilization process does not affect any additive contained in the coating.
  • the coating is mechanically durable and tough. It is unaffected by exposure to body fluids because of its highly inert character. Coatings containing elutable additives retain good durability following elution of additives including particulate additives.
  • the coating has good abrasion resistance for applications that may expose it to some degree of frictional wear. Further, the coating shows good resistance to tear propagation, even with high loading of additives.
  • FIG. 1A is a transverse cross section of an elongate article of round cross section such as a metal wire (for example, as from a self-expanding stent or an electrical conductor), or a polymeric suture, provided with a coating of the present invention.
  • a metal wire for example, as from a self-expanding stent or an electrical conductor
  • a polymeric suture provided with a coating of the present invention.
  • FIG. 1B is a transverse cross section of an article of rectangular cross section such as a stent element from a laser-cut balloon expandable stent, provided with a coating of the present invention.
  • FIG. 1C is a transverse cross section of the same article shown by FIG. 1B except that a partial coating of the present invention is provided, on only one surface of the article.
  • FIG. 1D is a transverse cross section of the same article shown by FIG. 1B except that first coating layer of the present invention is used that is provided with an additive, and then a second layer of the coating material is provided which does not contain an additive.
  • FIG. 1E is a transverse cross section of the same article shown by FIG. 1B except that two opposing sides of the article are provided with differently-filled coating layers.
  • FIG. 1F is a transverse cross section of the same article shown by FIG. 1B except that one surface of the article is provided with a first continuous layer of the inventive coating containing a first additive, and a second discontinuous layer of the coating material is provided containing a second additive different from the first additive.
  • FIG. 1G is a transverse cross section of the same article shown by FIG. 1B except that discontinuous layers of the coating material are provided on both surfaces of the article, leaving portions of the article surface exposed between the discontinuous segments of the coating.
  • FIG. 1H is a transverse cross section of the same article shown by FIG. 1B except that the article is provided with pockets that are filled with a first coating of the inventive material with a continuous second layer of the material being used as a cap over the first layer contained by the pockets in the article.
  • FIG. 1J is a cross section of multiple metallic stent elements provided with a continuous coating of the present invention that fully covers the stent elements and spans the interstices between the stent elements.
  • FIG. 2A is a perspective view of a laser-cut balloon expandable stent.
  • FIG. 2B is a top view of a section of the stent of FIG. 2A prior to deployment.
  • FIG. 2C is a top view of a section of the stent of FIG. 2A following deployment involving expansion of stent elements.
  • FIGS. 3A and 3B are perspective views of stent-grafts having a coating of the present invention.
  • FIG. 4A is a perspective view of a vascular graft provided with a coating of the present invention.
  • FIGS. 4B and 4C are transverse cross sections of coated vascular grafts of the present invention.
  • FIG. 4D is a perspective view of a helically-wrapped vascular graft of the present invention.
  • FIG. 5 is a longitudinal cross section of a catheter guidewire device or alternatively a helically wound electrical conductor provided with a coating of the present invention.
  • FIGS. 6A and 6B are isometric and cross sectional views of a keratoprosthesis using the coating material of the present invention.
  • FIG. 7A is a cross-sectional view of a composite two-layered device containing a dot-matrix pattern of a therapeutic agent applied between the two layers, shown as punctured by a needle.
  • FIG. 7B is a top view of the two-layered device of FIG. 7A, shown following removal of the puncturing needle.
  • FIG. 7C is a graph describing the release of drug from the device of FIGS. 7A and 7B, indicating a substantial, abrupt increase in the release rate following puncture with a needle.
  • FIG. 8A is a graph of normalized cumulative mass of dexamethasone released from three samples of wires provided with different polymer-drug coating formulations demonstrating a range of possible release kinetics.
  • the open circles depict the emulsion plus dispersion formulation, filled circles the single emulsion and filled triangles the powder coating formulation.
  • FIG. 8B is a scanning electron photomicrograph (SEM; about 20 ⁇ ) of a straight, 0.5 mm diameter wire provided with the coating of the present invention, and subsequently bent.
  • FIG. 9 is a scanning electron photomicrograph (about 260 ⁇ ) of a stent element provided with the coating of the present invention following a scrape test with a scalpel blade.
  • FIGS. 10A and 10B are scanning electron photomicrograph (about 100 ⁇ and 200 ⁇ , respectively) of two adjacent, curved expandable stent elements from balloon expandable stents, both provided with a coating of the present invention, steam sterilized and subsequently fully expanded.
  • FIGS. 11A and 11B are light micrographs (about 15 ⁇ ) a stent coated with three layers of drug/polymer, and subjected to ethylene oxide (EtO) sterilization at 67.7° C., before and after expansion with a balloon.
  • EtO ethylene oxide
  • FIGS. 12A and 12B are light micrographs (about 15X and 30X, respectively) of a TFE/PPVE polymer-coated stent that has been subjected to EtO sterilization at 67.7° C., before and after balloon expansion.
  • FIGS. 13A and 13B are light micrographs (about 15 ⁇ and 30 ⁇ , respectively) of a drug-TFE/PPVE polymer-coated stent that has been subjected to EtO sterilization at 67.7° C., before and after balloon expansion.
  • FIG. 14 is a light micrograph (about 10 ⁇ ) of a self-expanding stent-graft having a graft covering of the TFE/PMVE coating of the present invention.
  • FIGS. 15A and 15B are light micrographs (about 30 ⁇ ) of a TFE/PMVE polymer coated stent-graft, unsterilized, before and after balloon expansion.
  • FIGS. 16A and 16B are light micrographs (about 30 ⁇ and 15 ⁇ , respectively) of TFE/PMVE polymer coated stent-graft, unsterilized, before and after balloon expansion.
  • FIG. 17 is a graph of the cumulative mass of drug released as a function time and capping mass ratio for Example 14.
  • FIG. 18 is a graph of the cumulative mass of drug released as a function time for Example 15, where control device is filled triangles and test is open and filled circles.
  • FIGS. 19A and 19B are light micrographs (about 15 ⁇ and 20 ⁇ , respectively) of an EtO sterilized, TFE/PPVE-polymer coated stent-graft before and after balloon expansion.
  • FIGS. 20A and 20B are light micrographs (about 25 ⁇ and 30 ⁇ , respectively) of a TFE/PPVE-polymer coated stent-graft including a TFE/PMVE drug-containing layer, shown before and after expansion.
  • the present invention comprises a medical device in combination with a thermoplastic fluoropolymer, which is preferably an amorphous fluoropolymer.
  • the fluoropolymer may optionally contain various additives.
  • the thermoplastic fluoropolymer is a copolymer of tetrafluoroethylene (TFE) and perfluoroalkylvinylether (PAVE) that is free of cross-linking monomers and curing agents.
  • the perfluoroakylvinylether may be perfluoromethylvinylether (PMVE), perfluoroethylvinylether (PEVE) or perfluoropropylvinylether (PPVE).
  • the desirable mechanical characteristics, particularly tensile strength and toughness, are surprising given the absence of cross-linking monomers, curing agents, and process aids and fillers that would otherwise render such materials inadequately biocompatible.
  • the copolymer of TFE and PMVE is generally preferred, and may be made by emulsion polymerization techniques.
  • the PMVE content ranges from 40 to 80% by weight, while the complemental TFE content ranges from 60 to 20% by weight.
  • These materials have a secant modulus at 100% elongation of between 1 and 7 MPa (per ASTM D412-98, using 1 ⁇ 2 scale type IV dogbone with 250 mm/minute crosshead speed and 40 mm grip separation).
  • the material has a durometer in the range of 50-90 Shore A.
  • Durometer measurements are made at room temperature (about 23° C.) by the method of ASTM D2240-91 using a Shore Durometer Type O with a Shore model CV-71200 Conveloader (Shore Instrument Co., Freeport, N.Y.).
  • the durometer uses a hemispherical indenter of 1.2 mm radius. Samples tested by this method should be at least 6 mm thick; two or more samples may be stacked if necessary to achieve the minimum 6 mm thickness. Five durometer readings should be taken at five different points on each sample; these five readings are then averaged with the resulting mean value taken as the representative hardness value of the sample. Thickness measurements are the average of three or more measurements with a set of measuring calipers.
  • the PAVE component of the present invention is of the form
  • n the number of carbon atoms in the side chain, equals 1 to 3.
  • Copolymers of TFE/PAVE can be analyzed for copolymer composition with various characterization techniques known to those of skill in the art, including both nuclear magnetic resonance (NMR) spectroscopy and Fourier transform infrared (FTIR) spectroscopy, with NMR as the primary method, complemented and confirmed by FTIR.
  • NMR nuclear magnetic resonance
  • FTIR Fourier transform infrared
  • FIG. 1A is a transverse cross section of an elongate article 14 of round cross section such as a metal wire (for example, as from a self-expanding stent or an electrical conductor), or a polymeric suture, provided with a coating 12 of the present invention.
  • Coating 12 covers the entire surface of the article 14 to create a coated article 10 which may be of any shape.
  • Article 14 may be of any material other than the TFE/PAVE material of the coating.
  • Typical metallic materials for article 14 may be metals such as stainless steels, nitinol alloys, platinum, gold, silver, etc.
  • polymeric materials useful as article 14 include PTFE or ePTFE, polyethylene terephthalate (PET), polydimethylsiloxane (silicone), polyurethane (PU), or various other polymers known for use as medical devices. While the figure indicates that the entire outer surface of article 14 is provided with coating 12 , it is apparent that only selected portions of the surface of article 14 may be covered as desired.
  • coating 12 covers the entire surface (i.e., all surfaces) of article 14 , it is referred to as a continuous coating, that is, an uninterrupted coating that fully covers the article 14 .
  • Partial coatings that are interrupted in any of a variety of possible ways e.g., covering some surfaces while other surfaces remain uncovered, or dot-matrix pattern coatings, etc. are considered to be discontinuous coatings.
  • Coatings may be in single or multiple layers. Any layer can contain one or more additives such as therapeutic agents. Any of the layers may be provided in porous (e.g., containing void spaces) forms or non-porous forms.
  • FIG. 1B is a transverse cross section of an article 16 of rectangular cross section such as a stent element from a laser-cut balloon expandable stent, provided with a coating 12 of the present invention.
  • the article 16 may be made from a variety of materials and the coating 12 may be full or partial.
  • FIG. 1C is a transverse cross section of the same article 16 shown by FIG. 1B except that a partial coating 12 of the present invention is provided, on only one surface of the article.
  • FIG. 1D is a transverse cross section of the same article 16 shown by FIG. 1B except that first coating layer 12 a of the present invention is used that is provided with an additive, and then a second layer 12 b of the coating material is provided as a capping layer which does not contain an additive.
  • FIG. 1E is a transverse cross section of the same article 16 shown by FIG. 1B except that two opposing sides of the article are provided with differently-filled coating layers 12 c and 12 d.
  • FIG. 1F is a transverse cross section of the same article 16 shown by FIG. 1B except that one surface of the article is provided with a first continuous layer of the inventive coating 12 a containing a first additive, and a second discontinuous layer 12 e of the coating material is provided containing a second additive different from the first additive. It is apparent that discontinuous layer 12 e may be applied in any desired pattern, to any or all surfaces, etc. so that any desired pattern that is less than fully covering (i.e., continuous) may be produced.
  • FIG. 1G is a transverse cross section of the same article 16 shown by FIG. 1B except that a discontinuous layer 12 e of the coating material is provided on both surfaces of the article 16 , leaving portions of the article surface exposed between the discontinuous segments of the coating.
  • FIG. 1H is a transverse cross section of the same article 16 shown by FIG. 1B except that the article is provided with pockets 18 that are filled with a first coating 12 e containing an additive with a continuous second layer 12 b of the material being used as a cap over the first layer contained by the pockets in the article.
  • FIG. 1J is a cross section of multiple metallic stent elements of the present invention provided with a continuous coating that fully covers the stent elements and spans the interstices between the stent elements.
  • FIG. 2A is a perspective view of a laser-cut balloon expandable stent, intended as representative of stents generally.
  • Stent 22 is provided with a coating of the thermoplastic fluoropolymer. As stated previously, the coating may be continuous or discontinuous, and may be provided with a variety of additives.
  • the stent 22 is made from a suitable material such as any of various polymers or various metals including stainless steels or nitinols. While the stent shown is a balloon expandable stent, it is apparent that other types of stents including self-expanding stents may be coated as well.
  • Stent 22 is provided with a series of apices 24 that are plastically deformable during diametrical expansion of the stent.
  • FIG. 2B is an enlarged top view of a flattened section 22 a of the stent 22 of FIG. 2A prior to deployment.
  • Apices 24 have a relatively small radius prior to expansion.
  • FIG. 2C is an enlarged top view of the flattened section 22 a of FIG. 2B following deployment involving expansion of stent elements.
  • the previous relatively small radius of apices 24 is now increased due to plastic deformation resulting from stress applied during expansion. This deformation of stent apices 24 is problematic for many prior stent coatings in that they often crack or otherwise disrupt, with the result that the intended elution rate of any therapeutic agent contained in the coating can be significantly compromised.
  • Macroscopic cracking of the coating may be ascertained by expanding an endoluminal device under ambient conditions in an amount of 50 percent (measured as change in the outside diameter of the device) in accordance with the instructions for use for the particular device (if applicable), followed immediately by visual examination (aided if necessary by 10 ⁇ magnification).
  • the coating is typically unaffected by such a normal stent expansion, even when the coating is provided with a high additive content.
  • a device that is substantially free of such macroscopic cracks will have at most only a few minor cracks.
  • the capability of the coating of the present invention to be unaffected by deformation of stent components resulting from typical expansion can be demonstrated by providing a coating onto the surface of a wire.
  • the coating should be applied in a desired amount, loaded with the desired additive in the desired amount.
  • a straight length of wire having a round cross section of about 0.5 mm diameter should be used, with the wire being made of the same metal as a desired stent.
  • the wire is subjected to any sterilization procedure intended for the similarly coated stent. Following sterilization, the wire is bent at least 90 degrees at about the middle of its length, to a bend radius of 1.5 mm (i.e., to a bend radius of three times the wire diameter). The radius is measured to the inner meridian of the bent wire so that the wire can be bent around a form having a radius of 1.5 mm.
  • typically no cracking or other similar disruption of the coating will occur.
  • FIG. 3A describes a stent-graft 32 of the present invention wherein stent 22 is provided with a graft covering 34 .
  • the graft covering may be provided over the inner surface of the stent as described by FIG. 3A, or over the outer surface of the stent, or both the outer and inner surfaces of the stent.
  • Stent 22 may be any type of stent, including balloon expandable or self-expanding.
  • the stent 22 described by FIG. 3A is intended only to be representative of stent forms generally and as such is not intended as limiting.
  • the graft material may be made from a variety of materials including any known graft materials, for example, polyethylene terephthalate, silicone, polyurethane or ePTFE.
  • Stent-graft 32 is beneficially provided with a coating of the present invention that may optionally contain any of a variety of additives as described previously.
  • a stent-graft such as described by FIG. 3A may be provided with a continuous coating of the coating material, wherein the TFE/PAVE coating covers the stent elements and the graft covering material.
  • the entire graft covering may be coated including inner and outer surfaces. If the graft covering extends over only the inner or the outer surface of the stent (or any portion of those surfaces), the remaining surfaces of the stent that are not covered by the graft material may also be provided, or alternatively not provided, with the coating. Likewise, if desired, only the exposed portions of the stent 22 may be provided with the coating, leaving the graft material uncoated.
  • the coating may, for many inventive combinations of stent and graft materials, optionally be used as an adhesive to attach stent surfaces to the portions of the graft surfaces.
  • the stent-graft may be provided with a discontinuous coating 12 e according to the present invention.
  • This discontinuous coating can take a variety of forms as suggested by FIG. 3B.
  • a dot-matrix coating 12 e is applied over portions of the outer surface of the graft material covering the stent.
  • the dot-matrix coating may be provided with any of various additives in desired amounts. Different dots within the dot-matrix pattern may be provided with different therapeutic agents if desired.
  • a coating containing a first therapeutic agent may be provided to the luminal surface while another coating containing a second therapeutic agent different from the first may be applied to the exterior surface.
  • FIG. 4A describes a tubular vascular graft 42 provided with a coating of the present invention.
  • the coating may be continuous or discontinuous (including, for example, dot-matrix patterns) as described previously. Additives may be added to the coating as desired for any of a variety of purposes, also as described previously.
  • the vascular graft substrate material may be, for example, any known graft material such as ePTFE, PET or PU.
  • the coating 12 may be provided on the luminal surface of the graft substrate 44 .
  • the coating 12 may be provided as a middle layer between inner and outer layers of vascular graft substrate 44 .
  • the coating may be provided on the abluminal surface of the graft.
  • the coating may be impregnated into a portion or the entirety of the void space within the porous substrate.
  • FIG. 4D shows an ePTFE vascular graft substrate 44 provided with a helical wrap 46 of ePTFE film that has been provided as a narrow tape.
  • ePTFE films are made generally as taught by U.S. Pat. Nos. 3,953,566 and 4,187,390 to Gore.
  • the void space of the ePTFE film 46 may be impregnated with the coating described, or alternatively, the graft or the helically wrapped film may be coated as desired on any surface with the coating.
  • the helical wrap 46 may be in the form of the coating material.
  • the entire tubular vascular graft may be made from the coating material.
  • a vascular graft may be provided with a variety of additives as noted previously.
  • Such a graft may be formed with external mechanical support, such as molded in ridges, rings or struts. It is thus apparent that the coating may be applied in thicknesses as desired, to enhance the mechanical integrity or to provide other improved mechanical behavior to various medical devices in various ways. Coatings such as these may also incorporate additives.
  • FIG. 5 is a longitudinal, partial cross-section of a catheter guidewire device 52 or alternatively a helically wound electrical conductor 52 provided with a coating 12 .
  • Coating 12 may be provided continuously as shown or alternatively in a discontinuous form if desired; likewise the coating may be provided with one or more additives if desired.
  • the coating 12 may also be provided as a helical wrap of a tape made from the coating material.
  • FIG. 6A is an isometric view of an implantable cornea prosthesis or keratoprosthesis.
  • Keratoprosthesis 60 preferably having an ePTFE peripheral skirts or skirts 63 and 64 , is attached to a fluoropolymer cornea substitute 66 .
  • the skirts have a porosity that can be tailored to promote rapid ingrowth and attachment into surrounding native tissue.
  • FIG. 6B is a cross-sectional view of an implantable keratoprosthesis 60 , taken along section lines 62 , showing a first ePTFE skirt layer 63 , a second ePTFE skirt layer 64 and an polymeric cornea substitute layer 66 .
  • the cornea substitute layer 66 can be shaped to conform to surrounding native tissue and have a thickness, flexibility and creep resistance suitable for long term ocular implantation.
  • the ePTFE skirts can be pre-treated with a wetting agent such as poly(vinyl alcohol) to promote rapid post implant wetting, which enhances to initial anchoring to surrounding tissue.
  • Keratoprosthesis 60 can be produced, for example, by a lamination process in which one or more layers of ePTFE 63 , 64 are aligned onto a polymeric corneal layer 66 and compression molded to form a laminate.
  • polymeric corneal layer 66 can also be used to form an implantable lens or other light-transmitting device.
  • Additives such as ultraviolet absorbers, pigments or therapeutic agents can also be incorporated into the polymeric layer 66 , or into other optical devices such as lenses or transparent coatings.
  • a sample of TFE/PMVE copolymer was made by emulsion polymerization resulting in average emulsion particle size of 32 nanometers (particle size estimated using light scattering methods), exhibiting the following properties: mean tensile strength of 15.2 MPa, mean 100% secant modulus of 2.37 MPa, average tensile set of 0%, and PMVE content of about 66% by weight.
  • This copolymer sample was compression molded to produce a thin film of 0.18 mm thickness. Approximately 15 micrograms of the thin film in the form of a square sample of about 0.2 mm length per side was utilized for determination of the copolymer degradation temperature by themogravimetric analysis. The high-resolution scan covered the temperature range of 0-800° C. at heating rate of 20° C. per minute. Test results indicated that material degradation initiated at approximately 400° C., with a weight loss of less than about 0.5% at 400° C.
  • Thin films of TFE/PMVE copolymer described by Example 1 were produced via melt extrusion at temperatures exceeding 200° C.
  • a film possessing a thickness of approximately 0.2 mm was used to construct a laminate with pockets of chlorhexidine dihydrochloride, an antimicrobial agent.
  • a polypropylene template with 0.7 mm diameter holes arranged in a rectangle pattern was made to facilitate manufacturing of the device. The holes were evenly spaced approximately 2 mm apart, from edge to edge.
  • This template was placed on top of one of the TFE/PMVE extruded sheets, then dusted with chlorhexidine dihydrochloride. The template was removed, leaving a dot-matrix pattern of the drug on the surface of the extruded film.
  • a second sheet of extruded polymer was gently placed on top of the first sheet.
  • the composite of polymer sheets and drug was wrapped in aluminum foil, placed between two metal plates, heated in an oven set at 115° C. for 15 minutes, removed from the oven, immediately pressed between the two hot metal plates for 15 minutes, and then removed from the metal plates and aluminum foil. This process created encapsulated drug pockets between the polymer films.
  • the composite exhibited excellent bond characteristics.
  • the bond strength was so high that all attempts to delaminate the polymer films resulted in destruction of the composite.
  • FIG. 7A A cross-sectional view of the composite device 70 is shown in FIG. 7A.
  • First film layer 72 is provided as a cap over second film layer 74 , with film layer 74 being provided with a dot-matrix pattern 76 of a desired drug.
  • the device 70 is shown as it would appear when punctured with a needle 78 .
  • FIG. 7B illustrates device function following removal of the puncturing needle, allowing immediate release of drug from dots 76 that are affected by the needle puncture 79 .
  • the copolymer of Example 1 was obtained in a 4 wt % solution of FC-75.
  • the working drug formulation was a mixture of 2 ml of 4 wt % polymer, 8 ml of FC-75, and 150 mg of dexamethasone (52 wt % drug based on total weight of coating solids; dexamethasone obtained from Pharmacia & UpJohn, Kalamazoo Mich.).
  • the formulation was made by weighing dexamethasone into a test tube, adding FC-75, vortexing vigorously to complete mixing, adding the polymer, and ensuring complete mixing with additional vortexing.
  • a 10 cm length of Gore-Tex Vascular Graft (part number UT05070L, WL Gore & Associates, Flagstaff Ariz.) was used to demonstrate the drug release coating.
  • the 5 mm inside diameter graft was mounted onto a mandrel for coating.
  • the mandrel was rotated by hand as an airbrush (Badger standard set model 350 airbrush set at 220 KPa gauge air pressure, Badger Air Brush Co., Franklin Park, Ill.), held at a constant distance of approximately 3.8 cm from the graft surface, was moved back and forth across the graft while spraying a coating of the above-described polymer-drug formulation.
  • the vascular graft was continuously spray-coated for approximately 10 minutes, after which time the graft was transferred to an oven set at 60° C. for 2 minutes. Microscopic examination of cross sections of such a coated graft indicated that the coating penetrated into the void spaces of the microstructure of the porous ePTFE vascular graft. Physical examination of these coated graft samples indicated that the coating was well adherent.
  • the vascular graft was divided into two sections, 5 and 4 cm in length. A slight contraction of the graft in the longitudinal direction was noted after the coating was applied, as the total length measured about 9 cm after coating. This contraction was believed to be the result of drying of the relatively heavy coating.
  • the 5 cm section was coated with a capping layer that did not contain any drug.
  • the capping formulation consisted of 2 ml of 4 wt % polymer mixed with 8 ml of FC-75. The solution was sprayed in a similar manner as above in five 30 second spray intervals. Spraying intervals were separated by a 15 second interval of not spraying. The 4 cm section was sprayed in eight 30 second intervals, alternating with 15 second intervals of not spraying.
  • the 5 cm long section was taken for determination of total drug loading. Loading determinations were performed by placing the sample in 5 ml of ethanol in a glass test tube for 15 hours at 55° C. After ethanol extraction, the solution was analyzed for dexamethasone content using a UV spectrophotometer (dexamethasone wavelength: 244 nanometers). The loading was determined to be 7.5 ⁇ 1.0 mg/cm graft length.
  • TFE/PMVE not containing any drugs could also have been spray coated directly onto the vascular graft surface.
  • the coating may be applied between layers of the vascular graft, or may be applied to the luminal surface of a vascular graft.
  • a spray formulation was consisting of a dexamethasone emulsion plus dispersion was investigated first. Two ml of this 4 wt % polymer solution was diluted with 8 ml of FC-75 and mixed in a 15 ml plastic test tube, with periodic vortexing. 12.5 mg of dexamethasone as a powder and 200 microliters of a saturated ethanol solution containing dexamethasone (approximately 15 mg/ml of dexamethasone) were added to the solution. The system was vortexed for 1 minute to ensure complete mixing. It contained 10 wt % drug based on total weight of coatings solids, with wt % drug content calculated as drug mass/(drug+polymer mass), multiplied by 100.
  • the system was then coated onto a straight length of 0.51 mm diameter silver-plated copper wire.
  • This wire was intended to serve as a model of a structural element used in various medical devices, such as stents.
  • the wire was spray-coated with the polymer using a Badger standard set model 350 airbrush, for 1 minute, at an air pressure of 220 KPa, and placed in an air forced furnace for 5 minutes at 60° C. Ethanol extraction of such a coated wire segment followed by UV spectrophotometric analysis yielded 6.4 micrograms of drug per cm of wire length.
  • a capping layer that did not contain any drug was sprayed onto the wire.
  • the capping formulation consisted of 2 ml of 4 wt % polymer mixed with 8 ml of FC-75. The solution was sprayed in a similar manner as described above. The total coating on the wire was approximately 10 microns thick.
  • a spray formulation consisting of a single emulsion of dexamethasone was also investigated.
  • the working formulation was made by combining 2 ml of the 4 wt % polymer solution with 8 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • 400 microliters of a saturated ethanol solution containing dexamethasone (approximately 15 mg/ml of dexamethasone) was added to the copolymer solution.
  • the system was vortexed for 1 minute before coating to ensure complete mixing.
  • the coating on this wire was approximately 5 microns thick.
  • the coating contained 4.1 wt % drug based on total weight of coating solids. Ethanol extraction of wire segment followed by UV spectrophotometric analysis yielded 17.5 micrograms per cm of wire length.
  • a powder coating formulation was also investigated. Two ml of the 4 wt % polymer solution was combined with 8 ml of FC-75, then mixed in a 15 ml plastic test tube, with periodic vortexing. A polymer base coat was applied to the wire for 2 min. While still wet, the wire was suspended in a blender that had been pulsed briefly to air suspend dexamethasone. A capping layer that did not contain any drug was sprayed on the wire. The capping formulation consisted of 2 ml of 4 wt % polymer mixed with 8 ml of FC-75. The solution was sprayed in a similar manner as described above. The coating on this wire was approximately 5 microns thick. No theoretical loading was calculated.
  • FIG. 8A demonstrates the extended elution times possible with the different emulsion spray formulations, based on three samples made as described above (open circles depict the emulsion plus dispersion formulation, filled circles depict the single emulsion and filled triangles depict the powder coating formulation).
  • open circles depict the emulsion plus dispersion formulation
  • filled circles depict the single emulsion and filled triangles depict the powder coating formulation.
  • Each of the three types of coating resulted in smooth and uniform surfaces before and after drug release as evidenced by SEM analysis.
  • 8B is an SEM (about 20 ⁇ magnification) showing the crack-free mechanical integrity of the single emulsion coating of the different emulsion spray formulation process when the coated wire was bent in excess of a 90 degree angle, at a radius of about 1.1 mm as measured to the inside meridian of the bent wire.
  • the TFE/PMVE coating remained intact after complete elution of the drug.
  • TFE/PMVE copolymer made from emulsion polymerization resulting in an average emulsion particle size of about 120 nanometers, was prepared having the following properties: mean tensile strength of 26.7 MPa, mean 100% secant modulus of 2.7 MPa, mean tensile set of 12%, and PMVE content of about 60% by weight. Neither this TFE/PMVE copolymer nor any TFE/PMVE copolymer used in any the examples contained any cross-linking monomers or curing agents or system. The copolymer was added to FC-75 fluorinated solvent, to make a 4 wt % solution.
  • FC-75 fluorinated solvent 3M Fluorinert
  • the working formulation was made by diluting 2 ml of the 4 wt % polymer solution with 8 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • Stents made in accordance with the teachings of U.S. Pat. No. 5,925,061 to Ogi, et al. were laser cut and polished by Laserage Technology Corp., Waukegan, Ill. 60087. All stents were cut from 316H stainless steel tubing possessing a wall thickness of 0.127 mm. The outside diameter of the stents was 1.57 mm while the length was 21 mm.
  • Each stent was temporarily placed onto a small wire for handling during the coating process.
  • the wire was curved at one end to prevent the stent from slipping off.
  • the stent was dipped into the polymer solution, sprayed with compressed air to minimize any bridging of the coating between adjacent stent elements, and placed in an air forced furnace for 5 minutes at 60° C.
  • the dipping procedure may be repeated if multiple coatings are desired. For this example the dipping procedure was repeated 4 times. Scanning electron photomicrographs of uncoated and coated stents were taken before and after diametrically expanding up to 4.5 mm inner diameter with an angioplasty balloon. The expansion ratio was approximately 3.
  • FIG. 9 is a scanning electron photomicrograph (about 260 ⁇ magnification) of the surface after the scrape test. The coating was only removed from the regions of blade contact. There appeared to be no gross delamination or shrink-back of the coating from the scraped region, indicating good adhesion of the coating.
  • Coated stents made in this manner were steam sterilized (134° C. at 216 KPa for 12 minutes followed by a 30 minute drying cycle), balloon expanded to 3 mm diameter, and subjected to SEM analysis for determination of coating stability.
  • the scanning electron photomicrographs of FIG. 10A (about 100 ⁇ magnification) and FIG. 10B (about 200 ⁇ magnification) show that after processing and expansion, the polymer coating was still adherent to irregular shapes, without any evidence of delamination or tearing, demonstrating coating integrity even after steam sterilization and subsequent expansion.
  • a copolymer-drug coating where the TFE/PMVE copolymer is described by Example 1, was applied to balloon expandable stents of the same type as used in Example 5.
  • the amount of dexamethasone was approximately 400 micrograms per stent, applied by single emulsion spray coating as was done previously with the wire coating in Example 4.
  • the stent was balloon expanded to a diameter of 3.5 mm prior to initiating drug release studies. SEM analysis of the device surface subsequent to balloon expansion evidenced no delamination or separation of the coating from the metal. Release studies performed on another of these coated balloon expanded stents demonstrated that the drug was released in a controlled fashion. After completion of release studies, the sample underwent SEM analysis. The coating showed no delamination or separation from the metal.
  • the polymer-drug coating thickness was estimated to be approximately 3 microns.
  • stents were coated with the copolymer-drug solution through a dip coating processes in which the stents were suspended from a thin wire, immersed in the formulation, sprayed with compressed air at 1.7 KPa air pressure, and placed in a convection oven set at 60° C. briefly for compete drying.
  • One group of stents received 1 dip coating and another group 3 dip coatings.
  • Stents from each group were distended with the use of 3.5 mm PTFE balloons before and after sterilization with EtO at a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hours at 67.7° C.
  • Stents were examined with the use of a light microscope at magnification of up to 90 ⁇ . Microscopic examination of samples before and after expansion with or without EtO sterilization showed the coating to be tough, and well adhered, and without evidence of cracking.
  • FIG. 11A is a light micrograph (about 15 ⁇ magnification) of a drug/polymer coated stent according to this example that has been subjected to EtO sterilization at 67.7° C., before expansion. Three drug-polymer coat layers were applied to this stent as described above.
  • FIG. 11B (about 15 ⁇ magnification) describes the same stent after balloon expansion using a 3.5 mm diameter ePTFE/elastomer composite balloon (made generally as taught by example 7 of U.S. Pat. No. 6,120,477 to Campbell et al.). It is anticipated that virtually any suitable commercially available catheter balloon of suitable size would provide the same stent expansion results.
  • TFE/PPVE copolymer A sample of TFE/PPVE copolymer was obtained, which was synthesized by emulsion polymerization, resulting in average emulsion particle size of 83 nanometers, exhibiting the following properties: mean tensile strength of about 12.2 MPa, mean 100% secant modulus of 4.30 MPa, average tensile set of 31%, and PPVE content of about 56% by weight.
  • the polymer was dissolved in FC-75 to obtain a 20 wt % solution.
  • the working formulation was made by diluting 2 ml of the 20 wt % polymer solution with 8 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • Balloon expandable stents of the same type used in Example 5 were utilized. Each stent had a small wire temporarily looped through one end for handling during the subsequent dip-coating process. Once secured on the wire, the stent was dipped into the polymer solution, sprayed with compressed air, and placed in an air forced furnace for 5 minutes at 60° C. The dipping procedure was repeated to bring the total number of layers to 2. A portion of the TFE/PPVE coated stents were then expanded without being EtO sterilized using a balloon as described for Example 8, and examined with the use of a light microscope.
  • Additional coated stents underwent EtO sterilization with a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hours at 67.7° C. After sterilization the stents were expanded using a balloon of the type described for Example 8, and examined with a light microscope at magnification of up to 90 ⁇ . Microscopic examination of samples before and after expansion with or without EtO sterilization showed the coating to be tough, well adherent, and without evidence of cracking.
  • FIG. 12A is a light micrograph (about 15 ⁇ magnification) of a TFE/PPVE polymer coated stent according to this example that has been subjected to EtO sterilization at 67.7° C. before balloon expansion.
  • FIG. 12B is a light micrograph (about 30 ⁇ magnification) of the same stent following balloon expansion using a balloon as described in Example 8.
  • dexamethasone powder was weighed into a 15 ml plastic test tube, 6 ml of FC-75 was added, and the system was mixed vigorously to ensure complete mixing. Two hundred mg of 20 wt % TFE/PPVE polymer solution (made per Example 9) was added and the mixture was vortexed. This formulation is 60 wt % dexamethasone on a total solids basis. Balloon expandable stents of the same type used in Example 5 were utilized. Each stent had a small wire temporarily looped through one end for handling during the subsequent dip-coating process.
  • the stent was dipped into the polymer solution, sprayed with compressed air at 1.7 KPa air pressure, and placed in an air forced furnace for 5 minutes at 60° C. Stents were distended with the use of 3.5 mm PTFE balloons before and after sterilization with EtO at a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hours at 67.7° C.
  • Stents were examined with the use of a light microscope at magnification of up to 90 ⁇ . Microscopic examination of samples before and after expansion with or without EtO sterilization showed the coating to be tough, well adherent, and without evidence of cracking.
  • FIG. 13A is a light micrograph (about 15 ⁇ magnification) of a drug-TFE/PPVE-polymer coated stent made according to this example and subjected to EtO sterilization at 67.7° C., before balloon expansion.
  • FIG. 13B is a light micrograph (about 30 ⁇ magnification) of the same stent after balloon expansion using a balloon as described in Example 8.
  • the film initially contained void spaces, but these voids were filled as more layers were added.
  • This process can be practiced to produce a polymer stent cover that is perforated (i.e., containing occasional void spaces or openings through the coating that extends between adjacent wires) or continuous (without openings).
  • the polymer was dissolved in FC-75 to obtain a 4 wt % solution.
  • the working formulation was made by diluting 2.5 ml of the 4 wt % polymer solution with 5 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • Balloon expandable stents of the same type used in Example 5 were utilized. Each stent had a small wire temporarily looped through one end for handling during the subsequent dipcoating process. Once secured on the wire, the stent was dipped into the polymer solution, and placed in an air forced furnace for 5 minutes at 60° C.
  • FIG. 15A is a light micrograph (about 30 ⁇ magnification) of a TFE/PMVE polymer coated stent-graft according to this example shown before expansion while FIG. 15B is a light micrograph (about 30 ⁇ magnification) describing the same stent after balloon expansion using a balloon as described in Example 8.
  • the finished, coated stent-graft has occasional perforations or openings through the graft covering where substantial amounts of deformation of adjacent stent elements occurred during expansion.
  • FIG. 15A is a light micrograph (about 30 ⁇ magnification) of a TFE/PMVE polymer coated stent-graft according to this example shown before expansion
  • FIG. 15B is a light micrograph (about 30 ⁇ magnification) describing the same stent after balloon expansion using a balloon as described in Example 8.
  • the finished, coated stent-graft has occasional perforations or openings through the graft covering where substantial amounts of deformation of adjacent stent elements occurred during expansion.
  • 15B shows one such opening.
  • the more opaque regions of the coating adjacent to some stent elements were determined to be internal void spaces or “pockets” in the coating that were formed during stent expansion. They do not represent openings through the coating. While this is believed to be an artifact of the type of balloon-expandable stent used, it remains noteworthy that a large majority of the stent-graft covering was not occupied by these openings. For some applications, a stent-graft with occasional openings may be desirable. The stent-graft shown in this figure was not subjected to EtO sterilization.
  • the polymer was dissolved in FC-75 to obtain a 4 wt % solution.
  • the working formulation was made by diluting 3 ml of the 4 wt % polymer solution with 3 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • a Medi-tech 4 mm balloon (Boston Scientific, Medi-tech, Universal Product No. M001164180, Natick Mass.) was utilized to expand the stent-graft device.
  • FIG. 16A is a light micrograph (about 30 ⁇ magnification) of this TFE/PMVE polymer coated stent-graft before expansion.
  • FIG. 16B (about 15 ⁇ magnification) shows the same stent-graft immediately after balloon expansion to 4 mm.
  • the coating fully covers all of the stent interstices between adjacent stent elements, without any openings.
  • the more opaque regions of the coating adjacent to some stent elements were determined to be internal void spaces or “pockets” in the coating that were formed during stent expansion. They do not represent openings through the coating.
  • the stent-graft in this figure was not subjected to sterilization.
  • TFE/PMVE copolymer of Example 1 More of the same TFE/PMVE copolymer of Example 1 was obtained in a 2.5 wt % solution of FC-75.
  • the drug formulation was a mixture of 2 ml of 2.5 wt % polymer, 8 ml of FC-75, and 120 mg of dexamethasone. This solution was well-mixed by shaking and then sprayed with a Badger standard set model 350 airbrush set at 220 KPa gauge air pressure.
  • Porous expanded PTFE material was used to cover both the internal and external stent frame surfaces.
  • the inner ePTFE layer was constructed using an ePTFE tubing of about 25 microns thickness.
  • the outer surface of this inner layer was provided with a thin coating of the TFE/PMVE copolymer for subsequent use as a thermally-activated adhesive to join the ePTFE and stent layers.
  • the outer ePTFE layer was constructed by wrapping a 25 micron thick ePTFE tape about the outer stent surface. Both of these ePTFE materials were of about 25 micron average fibril length. These devices were placed into a convection oven set at 320° C. for five minutes to activate the adhesive. After removal from the oven and cooling to room temperature, the resulting 4 cm long stent-grafts were cut into three sections.
  • the scalloped end sections were cut to into 1.5 cm lengths and the mid-section was cut into a 1 cm length. Each of these sections was mounted onto a mandrel, rotated by hand and spray coated. The airbrush was held approximately 3.8 cm from the graft surface. Spraying was continuously performed for 30 seconds, after which time the coated stent-graft on the mandrel was transferred to an oven set at 60° C. for 2 minutes. This spraying and heating process was repeated for up to 21 times.
  • the devices were processed in three groups of 4 where, within each group, one stent-graft was for loading determination and the remaining 3 for release studies. The first group received 16 coats, the second 21, and the third 19 coats. Loading was periodically measured with the one stent-graft and the coating cycles adjusted to yield devices of comparable drug content.
  • a capping layer was applied with a solution of polymer made from 2 ml of the 2.5 wt % in 8 ml of FC-75. This was sprayed in a similar manner as was the drug containing formulation. Three groups consisting of three different capping layers were created by applying 5, 10 and 15 capping coats to the appropriate stent-graft group. The capping mass ratios are shown in FIG. 17.
  • FIG. 17 shows the cumulative mass of dexamethasone released as a function of time. Loading determinations were performed by placing the sample in 5 ml of ethanol in a glass test tube over night at 60° C. After ethanol extraction, the solution was analyzed by a UV spectrophotometer for dexamethasone content. Loading values for the 1.5 cm long stent-grafts were estimated to be 13.3, 12.8 and 15 mg for the respective groups. The capping mass was determined through gross weight change and determined to be 3.0, 6.0, and 8.5 mg, respectively.
  • stainless steel balloon expandable stents (about 1.5 mm unexpanded diameter) were obtained as described above.
  • the stent was powder-coated with FEP.
  • This ePTFE tube was placed over a mandrel, the powder-coated stent placed over the tube, and another ePTFE tube of the same type was placed over the stent.
  • the assembly was temporarily wrapped with an ePTFE film and placed in an oven set at 320° C. for five minutes.
  • the ePTFE tubes were thereby bonded to the stent, thereby encapsulating it and forming a stent-graft. After removal from the oven and cooling to room temperature, the temporarily-applied ePTFE film was removed.
  • TFE/PMVE copolymer made as described by Example 1 were utilized for coating the stent-graft. All formulations used polymer obtained in a 2.9 wt % solution of FC-75.
  • the first drug formulation was a mixture of 1 ml of 2.9 wt % polymer, 5 ml of FC-75, and 25 mg of dexamethasone. This solution was well mixed by vortexing and sprayed with a Badger standard set model 350 airbrush set at 220 KPa gauge pressure.
  • the stent-graft devices were placed onto mandrels and rotated by hand during the spraying process. The airbrush was held about 3.8 cm from the graft surface. In this manner only the abluminal surfaces of the devices were coated.
  • the second drug formulation was 1 ml of 2.9 wt % polymer, 5 ml of FC-75, 25 mg of dexamethasone, and 500 microliters of ethanol.
  • the system was mixed by sonication for 15 min. and vortexed briefly.
  • the third drug formulation was 1 ml of 2.9 wt % polymer, 5 ml of FC-75, 100 mg of dexamethasone, and 500 microliters of ethanol.
  • This solution was sprayed with a Badger, standard set model 350, spray paint gun set at 220 KPa gauge air pressure to coat devices.
  • Self-expanding stent-grafts of 15 mm length and 4, 4.5, and 5 mm diameters, of the type described in Example 14, were utilized. After the stent-grafts were mounted onto a mandrel, the mandrel was rotated by hand as the airbrush was moved back and forth across the stent-grafts. The airbrush was held at a constant distance of approximately 6 cm from the stent-graft surface. The coating was continuously sprayed for approximately 15 minutes, after which time the mandrel was transferred to an oven set at 60° C. for 2 minutes.
  • a capping layer was applied with a solution of polymer made from 2 ml of the 4 wt % in 8 ml of FC-75. This was sprayed for about 2.5 minutes, in a similar manner as the drug containing formulation, to obtain a capping mass of about 1.7 mg.
  • Each stent-graft was diametrically compacted to an outer diameter of approximately 2.1 mm; this may be accomplished by various means known to those of skill in the art of self-expanding stents.
  • the stent-grafts were constrained in the compacted state with a constraint wrap of more ePTFE film (not adhered), and were subjected to EtO sterilization with a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hrs at 54.4° C.
  • stent-graft devices were mounted onto a 3 mm angioplasty balloon, distended to the point of breaking the ePTFE film constraint wrap, and then fully distended with appropriate balloon sizes consistent with stent-graft diameters.
  • FIG. 18 shows the cumulative mass of dexamethasone released as a function of time for the control device (filled triangles) and test devices (open and filled circles). The control device was not compacted, sterilized, nor balloon distended; the test devices were subjected to all of these steps. The absence of spikes in the curves for the test grafts indicates the absence of cracking of the coating. Had the coating cracked, the drug elution curve would have demonstrated discontinuities associated with non-uniform delivery.
  • the two test stent-grafts show remarkable consistency in the release of dexamethasone after having been subjected to the physically challenging thermal and mechanical stresses. Furthermore, the test stent-grafts have retained the basic release characteristics of the control device with minimum deviation. From visual inspection of the curves in FIG. 18, it is evident that the curves are all very similar. From a pharmacokinetic standpoint two systems are generally equivalent if they deliver the same total quantity of drug and at the same rate (duration of delivery). The total drug delivered is take at the plateau regions of FIG. 18, and is determined to be 7.66 mg for control, and 6.935 mg and 6.811 mg for test samples. On a percentage basis the test samples are within 11 % of the control. This is remarkable in that the total drug loading for the devices is 10 mg, but only a consistent fraction of this is released as some remains trapped within the matrix. The test samples that underwent mechanical and thermal stress did not provide a total dose meaningfully different than the control.
  • TFE/PPVE copolymer described by Example 9 was prepared. The polymer was dissolved in FC-75 to obtain a 20 wt % solution. The working formulation was made by diluting 2 ml of the 20 wt % polymer solution with 8 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • Balloon expandable stents of the same type used in Example 5 were utilized. Each stent had a small wire temporarily looped through one end for handling during the subsequent dip-coating process. Once secured on the wire, the stent was dipped into the polymer solution, and placed in an air forced furnace for 5 minutes at 60° C. The dipping procedure was repeated to bring the total number of layers to 6. A portion of the stent-grafts were expanded before sterilization with a balloon as described in Example 8, and examined with the use of a light microscope. Additional coated stent-grafts underwent EtO sterilization with a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hours at 67.7° C.
  • the stent-graft was distended using a balloon as described in Example 8, and examined with a light microscope at magnification of up to 90 ⁇ .
  • the metal stent design can be utilized to produce a stent-graft having openings through the coating when expanded of predetermined size, and the metal stent design could be made to not facilitate the formation of openings resulting in an occlusive stent-graft post expansion.
  • FIG. 19A is a light micrograph (about 15 ⁇ magnification) of the TFE/PPVE-polymer coated stent-graft of this example shown before expansion, while FIG. 19B (about 20 ⁇ magnification) shows the same stent-graft after balloon expansion using a balloon as described in Example 8. Stent-grafts in this figure were EtO sterilized as described for previous examples.
  • a sample of the TFE/PPVE copolymer described by Example 9 was prepared.
  • the polymer was dissolved in FC-75 to obtain a 20 wt % solution.
  • the working formulation was made by diluting 2 ml of the 20 wt % polymer solution with 8 ml of FC-75 and allowing the system to mix in a 15 ml plastic test tube, with periodic vortexing.
  • a TFE/PMVE copolymer formulation containing the drug dexamethasone was also prepared.
  • the TFE/PMVE copolymer was dissolved in FC-75 to obtain a 4 wt % solution.
  • One hundred and twenty mg of dexamethasone as a powder was weighed into a 15 ml plastic test tube, 6 ml of FC-75 was added, and the system was mixed vigorously to ensure complete mixing.
  • Two grams of the 4 wt % TFE/PMVE polymer solution was added and the mixture was vortexed. This formulation is 60 wt % dexamethasone on a total solids basis.
  • Balloon expandable stents of the same type used in Example 5 were utilized. Each stent had a small wire temporarily looped through one end for handling during the subsequent dip-coating process. Once secured on the wire, the stent was dipped into the TFE/PPVE polymer solution, and placed in an air forced furnace for 5 minutes at 60° C. The dipping procedure was repeated to bring the total number of layers to 6. An additional layer containing the drug dexamethasone in TFE/PMVE was applied to the abluminal stent-graft surface. This was sprayed onto the stent-graft using a Badger, standard set model 350 airbrush set at 220 KPa gauge air pressure.
  • An end portion of the stent-graft was mounted onto a mandrel and then the mandrel was rotated by hand as the airbrush was moved back and forth across the stent-graft surface.
  • the coating was continuously sprayed for approximately 15 seconds, after which time the mandrel was transferred to an oven set at 60° C. for 2 minutes.
  • a portion of the stent-grafts were expanded with a balloon as described in Example 8, and examined with the use of a light microscope.
  • a coated stent-graft underwent EtO sterilization with a total cycle time of 15 hours, including an EtO sterilization time of 1.3 hours at 67.7° C. After sterilization the stent-graft was distended using a 3.5 mm PTFE balloon and examined with a light microscope at magnification of up to 90 ⁇ .
  • the drug-containing layer of TFE/PMVE did not separate from the base material of TFE/PPVE; and appeared to be tough, well adherent, and without evidence of cracking, demonstrating a high degree of stability. It is apparent that different copolymers of the PAVE family can be easily integrated into a single device construct, with or without additives.
  • FIG. 20A is a light micrograph (about 25 ⁇ magnification) of the TFE/PPVE-polymer coated stent-graft including the TFE/PMVE drug-containing layer, shown before expansion.
  • FIG. 20B (about 30 ⁇ magnification) shows the same stent-graft following expansion with a balloon of the type described in Example 8. While the covering shows occasional periodic and well-defined perforations or openings through the expanded stent-graft, the large majority of the stent-graft is unperforated. The stent-graft shown in these figures was not subjected to EtO sterilization.

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US10/213,126 2002-08-05 2002-08-05 Thermoplastic fluoropolymer-coated medical devices Abandoned US20040024448A1 (en)

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US10/213,126 US20040024448A1 (en) 2002-08-05 2002-08-05 Thermoplastic fluoropolymer-coated medical devices
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PCT/US2003/024618 WO2004012783A1 (en) 2002-08-05 2003-08-05 Thermoplastic fluoropolymer-coated medical devices
AU2003258112A AU2003258112B2 (en) 2002-08-05 2003-08-05 Thermoplastic fluoropolymer-coated medical devices
CA002493634A CA2493634C (en) 2002-08-05 2003-08-05 Thermoplastic fluoropolymer-coated medical devices
AT03767251T ATE423579T1 (de) 2002-08-05 2003-08-05 Thermoplastische fluoropolymer-beschichtete medizinprodukte
DE60326367T DE60326367D1 (de) 2002-08-05 2003-08-05 Thermoplastische fluoropolymer-beschichtete medizinprodukte
EP03767251A EP1545642B1 (en) 2002-08-05 2003-08-05 Thermoplastic fluoropolymer-coated medical devices
US11/378,401 US8048440B2 (en) 2002-08-05 2006-03-16 Thermoplastic fluoropolymer-coated medical devices
US13/282,892 US8609125B2 (en) 2002-08-05 2011-10-27 Thermoplastic fluoropolymer-coated medical devices
US14/107,343 US20140105955A1 (en) 2002-08-05 2013-12-16 Thermoplastic fluoropolymer-coated medical devices

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US8609125B2 (en) 2013-12-17
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US20140105955A1 (en) 2014-04-17
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CA2493634A1 (en) 2004-02-12
US8048440B2 (en) 2011-11-01
US20120065721A1 (en) 2012-03-15
AU2003258112A1 (en) 2004-02-23
AU2003258112B2 (en) 2008-02-07
CA2493634C (en) 2010-03-09
ATE423579T1 (de) 2009-03-15
EP1545642A1 (en) 2005-06-29
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EP1545642B1 (en) 2009-02-25
US20060198866A1 (en) 2006-09-07

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