WO2013158619A2 - Implant médical d'élution de médicament - Google Patents

Implant médical d'élution de médicament Download PDF

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Publication number
WO2013158619A2
WO2013158619A2 PCT/US2013/036743 US2013036743W WO2013158619A2 WO 2013158619 A2 WO2013158619 A2 WO 2013158619A2 US 2013036743 W US2013036743 W US 2013036743W WO 2013158619 A2 WO2013158619 A2 WO 2013158619A2
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WO
WIPO (PCT)
Prior art keywords
implant
coating
implants
medical implant
diameter
Prior art date
Application number
PCT/US2013/036743
Other languages
English (en)
Inventor
Maria Palasis
Changcheng You
Daniel Concagh
Lee Core
Kicherl Ho
Upma Sharma
Greg Zugates
Original Assignee
480 Biomedical, Inc.
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Priority claimed from US13/766,294 external-priority patent/US8992601B2/en
Application filed by 480 Biomedical, Inc. filed Critical 480 Biomedical, Inc.
Publication of WO2013158619A2 publication Critical patent/WO2013158619A2/fr

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Classifications

    • 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
    • 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
    • A61L31/14—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L31/16—Biologically active materials, e.g. therapeutic substances
    • D—TEXTILES; PAPER
    • D04—BRAIDING; LACE-MAKING; KNITTING; TRIMMINGS; NON-WOVEN FABRICS
    • D04C—BRAIDING OR MANUFACTURE OF LACE, INCLUDING BOBBIN-NET OR CARBONISED LACE; BRAIDING MACHINES; BRAID; LACE
    • D04C1/00—Braid or lace, e.g. pillow-lace; Processes for the manufacture thereof
    • D04C1/06—Braid or lace serving particular purposes
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
    • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
    • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
    • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
    • A61F2250/00—Special features of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof
    • A61F2250/0014—Special features of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof having different values of a given property or geometrical feature, e.g. mechanical property or material property, at different locations within the same prosthesis
    • A61F2250/003—Special features of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof having different values of a given property or geometrical feature, e.g. mechanical property or material property, at different locations within the same prosthesis differing in adsorbability or resorbability, i.e. in adsorption or resorption time
    • 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
    • A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
    • A61L2300/40—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices characterised by a specific therapeutic activity or mode of action
    • A61L2300/416—Anti-neoplastic or anti-proliferative or anti-restenosis or anti-angiogenic agents, e.g. paclitaxel, sirolimus
    • 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
    • A61L2420/00—Materials or methods for coatings medical devices
    • A61L2420/08—Coatings comprising two or more layers
    • D—TEXTILES; PAPER
    • D10—INDEXING SCHEME ASSOCIATED WITH SUBLASSES OF SECTION D, RELATING TO TEXTILES
    • D10B—INDEXING SCHEME ASSOCIATED WITH SUBLASSES OF SECTION D, RELATING TO TEXTILES
    • D10B2509/00—Medical; Hygiene
    • D10B2509/06—Vascular grafts; stents

Definitions

  • the present invention relates to medical implants, and more specifically, to self- expanding medical implants that are intended for placement within a lumen or cavity of a patient.
  • tubular medical implants include woven grafts and stent-grafts that are used to span vascular aneurysms, polymeric tubes and catheters that are used to bypass strictures in the ureter and urethra, and stents that are used in the peripheral vasculature, prostate, sinus and esophagus.
  • stents for the esophagus and other large bodily lumens may be as large as 30mm or 40mm in diameter or larger.
  • biodegradable and biostable polymeric materials for stents and other implantable devices has been proposed to eliminate the possible long-term effects of permanent implants, the use of such materials has been hindered by relatively poor expandability and mechanical properties.
  • the expansion characteristics and radial strength of prototype stents made from biodegradable and biostable polymeric materials has been significantly lower than that of metallic stents. This is particularly the case where such stents are low profile and make use of small diameter fibers or thin walled struts that comprise the stent body.
  • the degradation rate and the manner in which such devices degrade in the body have been difficult to control.
  • the drug elution rate has been difficult to reproducibly characterize.
  • the present invention includes an implantable medical device for placement within a lumen or cavity of a patient.
  • the present invention includes a method of loading the medical device into a delivery catheter just prior to being implanted into a patient.
  • the present invention includes a method of treating a patient by delivering the medical device to a target location within the patient.
  • the present invention includes a kit that comprises the implantable medical device.
  • the implantable medical devices of the present invention are generally tubular, self- expanding devices.
  • the devices have a combination of structure, composition, and/or strengthening means that provide them with exceptional expandability and mechanical properties when compared with conventional bioresorbable stents or self-expanding devices.
  • the implantable medical device comprises a tubular structure comprising a first polymeric material, a first conformal coating comprising poly (lactic acid-co-caprolactone) and at least partially coating the tubular structure, and a second conformal coating comprising a second polymeric material and paclitaxel.
  • the first conformal coating increases the mass of the tubular structure by at least about 20 percent.
  • the second conformal coating comprises between 0.3 weight percent and 5 weight percent of paclitaxel.
  • the second conformal coating comprises between 2.5 micrograms and 20 micrograms of paclitaxel per 10mm of length of the tubular structure.
  • FIG. 3 is a cross-sectional view of a strand of an implantable medical device in accordance with an embodiment of the present invention that includes a support coating.
  • Fig. 4 is a side view of a strand of an implantable medical device in accordance with an embodiment of the present invention that includes discrete areas of therapeutic agent thereon.
  • Fig. 6 is an end view of an implantable medical device in accordance with an embodiment of the present invention.
  • Fig. 7 is a graph of diameter recovery as a function of support coating weight, for certain embodiments of the present invention.
  • Fig. 8 is a graph of diameter recovery as a function of gel content within a support coating, for certain embodiments of the present invention.
  • Figs. 9 and 10 illustrate the effect of chain terminators on the acute RRF and COF of thermoset elastomer-coated stents, in accordance with certain embodiments of the present invention.
  • Fig. 12 illustrates the effect of DD on the COF of PLCL/HDI/DD-coated devices over time, in accordance with an embodiment of the present invention.
  • Fig. 13 illustrates the effect of fiber diameter on the RRF and COF of coated devices of the present invention.
  • Fig. 15 is a graph of time-dependent COF values of certain embodiments of the present invention, as well as for known commercial self-expanding stents.
  • the present invention provides for medical implants that have expansion characteristics and mechanical properties that render them suitable for a broad range of applications involving placement within bodily lumens or cavities.
  • devices that have expansion characteristics and mechanical properties that render them suitable for a broad range of applications involving placement within bodily lumens or cavities.
  • implant As used herein, “device,” “implant,” and “stent” are used synonymously.
  • self-expanding is intended to include devices that are crimped to a reduced configuration for delivery into a bodily lumen or cavity, and thereafter tend to expand to a larger suitable configuration once released from the delivery configuration, either without the aid of any additional expansion devices or with the partial aid of balloon-assisted or similarly-assisted expansion.
  • the implants of the present invention When compared with conventional self-expanding medical implants, the implants of the present invention recover to an exceptionally high percentage of their manufactured diameter after being crimped and held in a small diameter for delivery into a bodily lumen. Moreover, when compared with conventional self-expanding implants and particularly polymeric implants, the implants of the present invention are characterized by much improved strength and other desired mechanical properties. As used herein, “strength” and “stiffness" are used
  • the implant 100 preferably comprises at least one strand woven together to form a substantially tubular configuration having a longitudinal dimension 130, a radial dimension 131, and first and second ends 132, 133 along the longitudinal dimension.
  • woven is used synonymously with "braided.”
  • the tubular configuration may be woven to form a tubular structure comprising two sets of strands 1 10 and 120, with each set extending in an opposed helix configuration along the longitudinal dimension of the implant.
  • the sets of strands 1 10 and 120 cross each other at a braid angle 140, which may be constant or may change along the longitudinal dimension of the implant.
  • the braid angle 140 is within the range of about 90 degrees to about 135 degrees throughout the implant.
  • the strands are woven together using methods known in the art, using known weave patterns such as Regular pattern "1 wire, 2-over/2-under", Diamond half load pattern "1 wire, 1 -over/ 1 -under”, or Diamond pattern "2 wire, 1 -over/1 -under”.
  • the strands may be made from biostable polymeric or metallic materials, they are preferably made from at least one biodegradable material that is preferably fully absorbed within about two years of placement within a patient, and more preferably within about one year of placement within a patient. In some embodiments, the strands are fully absorbed within about six or fewer months of placement within a patient.
  • the first and second strand sets 1 10, 120 may be made from the same or different biodegradable polymer.
  • Non- limiting examples of biodegradable polymers that are useful in the at least one strand of the present invention include poly lactic acid (PLA), poly glycolic acid (PGA), poly trimethylene carbonate (PTMC), poly caprolactone (PCL), poly dioxanone (PDO), and copolymers thereof.
  • Preferred polymers are poly(lactic acid co-glycolic acid) (PLGA) having a weight percentage of up to about 20% lactic acid, or greater than about 75% lactic acid (preferably PLGA 85: 15), with the former being stronger but degrading in the body faster.
  • the composition of PLGA polymers within these ranges may be optimized to meet the mechanical property and degradation requirements of the specific application for which the implant is used.
  • the materials used for the strands preferably have an elastic modulus within the range of about 1 to about 10 GPa, and more preferably within the range of about 6-10 GPa.
  • the strands used in the implant 100 preferably have a diameter in the range of from about 125 microns to about 225 microns, and are more preferably less than about 150 microns in diameter.
  • the use of small diameter strands results in an implant with minimal wall thickness and the preferred ability to collapse (i.e., to be crimped) within low diameter catheter delivery systems.
  • multiple strands may be of substantially equal diameters within this range, or first strand set 1 10 may be of a different general diameter than second strand set 120.
  • the diameters of strands are chosen so as to render the implant 100 preferably deliverable from a 10 French delivery catheter (i.e., 3.3mm diameter) or smaller, and more preferably from a 7 French delivery catheter (i.e., 2.3 mm diameter) or smaller.
  • the ability to place the implant of the present invention into small diameter delivery catheters allows for its implantation into small diameter bodily lumens and cavities, such as those found in the vascular, biliary, uro-genital, iliac, and tracheal-bronchial anatomy.
  • Exemplary vascular applications include coronary as well as peripheral vascular placement, such as in the superficial femoral artery (SFA). It should be appreciated, however, that the implants of the present invention are equally applicable to implantation into larger bodily lumens, such as those found in the gastrointestinal tract, for applications such as esophageal scaffolds.
  • the implant 100 When the implant 100 is a unitary framework, it is fabricated using any suitable technique, such as by laser cutting a pattern into a solid polymer tube. In a preferred embodiment, when the implant 100 is a unitary framework, it is formed by laser cutting and includes a wall thickness of between about 75 and about 100 microns. It should be recognized that while the present invention is described primarily with reference to woven strand configurations, aspects of the present invention are equally applicable to non- woven, self-expanding structures unless necessarily or expressly limited to woven
  • RRF radial resistive force
  • COF chronic outward force
  • the self-expanding implants of the present invention should preferably recover to a high percentage of their as-manufactured configuration after being crimped for insertion into the body, and the thus expanded implant should have a relatively high RRF to be able to hold open bodily lumens and the like, yet have a relatively low COF so as to avoid applying possibly injurious forces against the walls of bodily lumens or the like.
  • the implants of the present invention preferably expand to at least 90% of their as manufactured configuration after being crimped, have an RRF of at least about 200mm Hg and preferably greater than 400mm Hg, have an acute COF (at the time of delivery into a bodily lumen or cavity) of about 40-200mm Hg, and a COF of less than about 10mm Hg after about 28 days in vivo.
  • the support coating 410 is made from a shape memory polymer or a polymer that otherwise contracts upon heating to body temperature.
  • the inventors have surprisingly found that the use of support coatings on the polymeric implants of the present invention can result in the recovery of more than 90% of implant diameter post- crimping, and yield significantly higher radial forces when compared with uncoated implants or even with self-expanding metallic stents.
  • the support coating 410 may be applied as a conformal coating (as shown in cross-section of an individual strand in Fig. 3), may be partially applied to one or more individual strands such that the support coating 410 is applied to only a portion of the implant along its longitudinal dimension, or may be applied to only the inner or outer diameter of one or more individual strands.
  • the support coating 410 may optionally vary in weight along the length of the implant; for example, the ends of the implant may be coated with a thicker layer than the mid-section to provide added support to the ends of the implant. In addition, the support coating may accumulate at the crossover points or "nodes" of the woven device, which has the effect of aiding in diameter recovery and the achievement of preferred COF and RRF values.
  • polymer materials used for the support coating 410 include suitable thermoplastic or thermoset elastomeric materials that yield the elongation, mechanical strength and low permanent deformation properties when combined with the implant strand(s).
  • suitable polymers include certain random copolymers such as poly(lactic acid-co-caprolactone) (PLCL), poly(glycolide-co-caprolactone) (PGCL), and poly(lactic acid-co-dioxanone) (PLDO), poly(butylene succinate) (PBS), poly(p-dioxanone) (PDO), certain homopolymers such as poly trimethylene carbonate (PTMC), and copolymers and terpolymers thereof.
  • PLCL poly(lactic acid-co-caprolactone)
  • PGCL poly(glycolide-co-caprolactone)
  • PLDO poly(lactic acid-co-dioxanone)
  • PBS butylene succinate
  • PDO poly(p-dioxanone)
  • certain homopolymers such as poly trimethylene carbonate (PTMC), and copolymers and terpolymers thereof.
  • such polymers are optionally crosslinked with a crosslinker that is bi- or multi-functional, polymeric, or small molecule to yield a thermoset polymer having a glass transition temperature (Tg) that is preferably lower than body temperature (37°C), more preferably lower than room temperature (25°C), and most preferably lower than about 10°C.
  • Tg glass transition temperature
  • the thermoset elastomers provide a high elongation to break with low permanent deformation under cyclic mechanical testing.
  • the polymer material used for the support coating 410 is a biodegradable thermoset elastomer synthesized from a four arm PGCL polymer having a weight ratio of approximately 50:50 GA:CL that is crosslinked with hexamethylene diisocyanate (HDI) to give a polyester with urethane crosslinks.
  • HDI hexamethylene diisocyanate
  • the inventors believe that the combination of the elastic segment (polyester portion) and the interactions (such as hydrogen bonding, allophanate or biuret formation) between the urethane segments of such polymers, in addition to a certain crosslinking density, yields preferred properties such as a high degree of elastic recovery under cyclic mechanical strain and high overall elasticity.
  • the support coating comprises PLCL having a weight ratio of approximately 50:50 LA:CL.
  • a PLCL 50:50 crosslinked with hexamethylene diisocyanate support coating is applied to a PLGA 75:25 braided implant.
  • the polymer material used for support coating 410 may be optimized for desired mechanical properties.
  • the inventors have found that the molecular weight of such polymers may be manipulated to enhance coating performance.
  • a molecular weight (Mn) between about 30kDa - el and 120kDa, preferably from 33k to 80k, results in a lower modulus of elasticity and a higher strain to fracture, thus making the coating better able to adhere to a PLGA braided implant during crimping and post-crimping expansion and therefore less likely to fracture during deployment.
  • a molecular weight (Mn) from 8kDa to 20kDa does not yield an appreciable change in properties, but that a further increase in molecular weight to 50kDa results in a four-fold increase in the strain at failure of the coating material.
  • Mn molecular weight
  • a preferred range of molecular weight (Mn) for PGCL used in the implants of the present invention is about 23kDa to about lOOkDa.
  • the inventors have found that the viscosity of the spray coating solution, the weight percent of crosslinker used in the spray coating solution, and the temperature and duration of the support coating crosslinking process can be optimized to provide preferred coating morphologies and radial forces.
  • the crosslink density may be varied to yield desired mechanical properties.
  • chain terminators may be used in thermoset elastomeric materials such as polyester urethanes to control crosslink density.
  • the chemical crosslink density is adjusted by using such chain terminators to control the degree of crosslinking taking place during the polyester-urethane curing.
  • the crosslink density of the resultant elastomers depends on the concentration of chain terminators incorporated into the elastomeric network. Both small molecular agents and polymers may react with chain terminators.
  • HDI is used as a cross- linker, and the prepolymer is provided in a ratio of between 1 : 1 and 25: 1 wt/wt relative to HDI.
  • Any suitable unreactive organic solvent may be used in the present invention, including dichloromethane (DCM), ethyl acetate, acetone, methyl tert-butyl ether, toluene, or 2- methyltetrahydrofuran.
  • suitable chain terminators include small molecular agents and polymers that carry one active functional group that may react with either isocyanate or hydroxyl groups, such as but not limited to amines, alcohols, isocyanates, acyl chlorides, and sulfonyl chlorides.
  • the suitable chain terminator is provided in a wide range of amounts (0- 20wt% relative to the prepolymer) to control the cross-linking density.
  • the solution of polyester, crosslinker, and chain terminator is dissolved in a solvent and spray coated onto the surface of the implant 100 and cured to form support coating 410 as a conformal elastomeric coating.
  • the support coating 410 is coated onto the surface of the implant 100 using any suitable method, such as spraying, dipping, electrospraying, rolling, and the like. If implant 100 is a woven structure, the support coating 410 may be applied to individual strands prior to forming the woven structure, or to the woven structure after the formation thereof. In this case, owing to surface tension, the coating preferably collects at intersection points between strands. If implant 100 is a non- woven structure, the support coating 410 may be applied, for example, to a solid polymer tube either before or after the removal of material such as by laser cutting to form a patterned, non-woven structure.
  • the amount of support coating 410 applied to the implant 100 has been identified as one of the factors that contribute to the expansion characteristics and mechanical strength of the implant.
  • the application of the support coating 410 increases the weight of the uncoated implant 100 by about 20% to about 100%, more preferably, by about 24% to about 70%, and most preferably by about 30% to about 60%.
  • the strengthening means includes the attachment of strand sets 110, 120 at one or more intersections of the strands along the longitudinal dimension of the implant 100.
  • Such attachment may be achieved by use of an adhesive, or by fusing the strands at predetermined intersections such as by heat, laser, or ultrasound.
  • the strands comprise materials having sufficient elasticity such that the local strains induced by the adhesive/welded joints would not cause plastic deformation thereof.
  • the strengthening means includes the incorporation of additives into one or more of the strands.
  • additives are neutralizing agents such as calcium salts (e.g., calcium carbonate or calcium phosphate) or other salts such as barium salts that increase the mechanical strength of the strands into which they are incorporated, and further act to neutralize any acidic byproducts resulting from the degradation of the strand material(s).
  • such additives are plasticizers such as polyethylene glycol (PEG) that dissolve from the strand(s) in vivo, thus increasing the flexibility of the strand(s) and the implant over time.
  • PEG polyethylene glycol
  • the implant 100 delivers one or more therapeutic agents at the site of implantation.
  • the therapeutic agent(s) may be applied to one or more strands for delivery therefrom in a number of ways.
  • the therapeutic agent(s) are embedded within a conformal polymer coating 210 that adheres to one or more individual strands of the implant 100.
  • a coating 210 is preferably made from a biodegradable polymer admixed with the therapeutic agent(s) such that the agent is eluted from the polymer over time, or is released from the coating as it degrades in vivo.
  • one or more therapeutic agents are applied in discrete areas 220 on one or more individual strands (shown as length of individual strand).
  • discrete areas 220 are preferably made from a biodegradable polymer admixed with the therapeutic agent(s) and eluted from the polymer over time, or are released from the coating as it degrades in vivo.
  • the biodegradable polymer may be the same as or different from the biodegradable polymer(s) used in the strands of the implant.
  • the therapeutic agent(s) are admixed within the strand(s) of the implant 100 such that the agent(s) are eluted from the one or more strands over time, or are released from the one or more strands as the strand(s) degrade in vivo.
  • the therapeutic agent(s) are admixed within a support coating, as described herein.
  • the therapeutic agent(s) may be admixed with the polymer used to fabricate the implant 100.
  • the implant designs of the present invention may present a significantly higher surface area, with more closely spaced strands or struts, when compared with conventional metallic stents, thus offering the possibility of a more uniform distribution of drug in the tissue surrounding the implant.
  • the therapeutic agent(s) used in the present invention are any suitable agents having desired biological effects.
  • the therapeutic agent is selected from anti-thrombogenic agents such as heparin, heparin derivatives, urokinase, and PPack (dextrophenylalanine proline arginine chloromethylketone), enoxaparin, hirudin; anti-proliferative agents such as angiopeptin, or monoclonal antibodies capable of blocking smooth muscle cell proliferation, acetylsalicylic acid, paclitaxel, sirolimus, tacrolimus, everolimus, zotarolimus, vincristine, sprycel, amlodipine and doxazosin; anti-inflammatory agents such as glucocorticoids, betamethasone,
  • immunosuppressants such as sirolimus, tacrolimus, everolimus, zotarolimus, and dexamethasone
  • antineoplastic/antiproliferative/anti- mitotic agents such as paclitaxel, 5-fluorouracil, cisplatin, vinblastine, cladribine, vincristine, epothilones, methotrexate, azathioprine, halofuginone, adriamycin, actinomycin and mutamycin
  • anesthetic agents such as lidocaine, bupivacaine, and ropivacaine
  • anti-coagulants such as D-Phe-Pro-Arg chloromethyl ketone, an RGD peptide-containing compound, heparin, antithrombin compounds, platelet
  • Preferred therapeutic agents used in the present invention to treat restenosis and similar medical conditions include sirolimus, everolimus, zotarolimus, vincristine, sprycel, dexamethasone, paclitaxel, and analogs thereof. Also preferred is the use of agents that have a primary mechanism of action of inhibiting
  • Coating 210 or areas 220 containing one or more therapeutic agents are applied to implant 100 by any suitable method, such as spraying, electrospraying, rolling, dipping, chemical vapor deposition, and potting.
  • coating 210 or areas 220 is further coated with a biodegradable or biostable topcoat as shown in Fig. 5 (individual strand shown in cross-section), that acts to regulate the delivery of the therapeutic agent from coating 210 or areas 220 into bodily tissue.
  • the topcoat 211 acts as a diffusion barrier such that the rate of delivery of the therapeutic agent(s) are limited by the rate of its diffusion through the topcoat 21 1. In another embodiment, the therapeutic agent(s) cannot diffuse through the topcoat 211 such that delivery thereof is simply delayed until the degradation of the topcoat 211 is complete.
  • the topcoat 211 preferably comprises a biodegradable polymer that is the same as or different from that of the coating 210 or the strands. If implant 100 is a woven structure, coatings 210, 220, or 211 may be applied to individual strands prior to forming into the woven structure, or to the woven structure after the formation thereof.
  • coatings 210, 220, or 211 may be applied, for example, to a solid polymer tube either before or after the removal of material such as by laser cutting to form a patterned, non-woven structure.
  • the coatings 210, 220, and/or 211 are preferably applied over such support coating 410, or the support coating 410 itself may include the therapeutic agent(s).
  • the coatings 210, 220, and/or 21 1 may be applied between the supporting coating 410 and the implant 100.
  • the implant 100 of the present invention is preferably self-expanding in that it is manufactured at a first diameter, is subsequently reduced or "crimped" to a second, reduced diameter for placement within a delivery catheter, and self-expands towards the first diameter when extruded from the delivery catheter at an implantation site.
  • the first diameter is preferably at least 20% larger than the diameter of the bodily lumen into which it is implanted.
  • the implant 100 is preferably designed to recover at least about 80% and up to about 100% of its manufactured, first diameter.
  • implants in accordance with the present invention have a recovery diameter greater than 80%, and preferably greater than 90%, of the manufactured, first diameter after being crimped into exemplary delivery catheters of either 1.8mm or 2.5mm inner diameter and held for one hour at either room temperature (25°C) or body temperature (37°C).
  • Such implants a braided structure comprising 32 strands ranging in size from 0.122mm to 0.178mm, braided with an inner diameter ranging from 5 to 6mm.
  • the implant 100 has a manufactured, first diameter (inner) of about 4.0mm to about 10.0mm.
  • the woven implants of the present invention have a braid angle 140 within a range of about 90° to 135°, more preferably within a range of about 1 10° to 135°, and most preferably within a range of about 115° to 130°.
  • the inventors have confirmed through experimentation that braid angle may affect certain mechanical properties of the implants of the present invention.
  • the inventors have found that while a braid angle of 110° for 6mm PGCL coated braided implants made from PLGA 10:90 copolymer strands yields post-crimp RRF and COF values at 4.5mm of 370mm and 147mm Hg, respectively, the same coated implants having a 127° braid angle are characterized by post- crimp RRF and COF values at 4.5mm of 900mm and 170mm Hg, respectively.
  • the ends of strands are fused in patterns on the outside and/or inside of the implant 100, as shown in Fig. 6. Fusing the strands in this manner causes the ends 132, 133 to flare outwards instead of curling inward, thus addressing a phenomenon in which the ends of the implant may naturally curl inward when the implant is crimped into a reduced configuration for insertion into a bodily lumen or cavity.
  • the ends are fused in patterns such that their final positions are axially staggered. Strands are fused, for example, by using heat, chemicals, adhesives, crimp swages, coatings (elastomeric or non-elastomeric), or other suitable fusing or joining techniques.
  • one or more biostable or biodegradable radiopaque markers preferably comprising platinum, iridium, tantalum, and/or palladium are produced in the form of a tube, coil, sphere, or disk, which is then slid over one or more strands of fiber to attach to the ends of implant 100 or at other predetermined locations thereon.
  • the marker is in the form of a tube or coil, it has a preferable wall thickness of about 0.050 to 0.075mm and a length of about 0.3 to 1.3mm.
  • the tube is formed by extrusion or other methods known in the art.
  • the coil is formed by winding a wire around a mandrel of desired diameter and setting the coil with heat or other methods known in the art.
  • the implant 100 is preferably loaded into a delivery catheter just prior to being implanted into a patient. Loading the implant 100 in close temporal proximity to implantation avoids the possibility that the polymer of the implant 100 will relax during shipping, storage, and the like within the delivery catheter and therefore cannot fully expand to a working configuration.
  • one aspect of the invention includes a method of delivering an implant of the invention that comprises the step of loading the implant into a delivery catheter within a short period of time, and preferably within one hour, before implantation into a body lumen. It should be noted, however, that it is not required that the implants of the present invention are loaded into delivery catheters just prior to being implanted. In fact, one advantage of the present invention is that it provides self-expanding implantable medical devices with preferred expansion characteristics and mechanical properties even after being loaded in a delivery catheter for prolonged periods.
  • Braided implants were manufactured using a PLGA 10:90 copolymer by spooling fiber spun monofilaments onto individual bobbins. Each bobbin was placed on a braiding machine, strung through rollers and eyelets and wrapped around a mandrel. The braiding tension of the machine was set for the size of the monofilament (i.e., 70g min, 215g max for 0.005" fiber). The pix/inch was set to obtain an ideal angle to maximize radial strength but still allow the braid to be removed from the mandrel (i.e., 110 to 135 degrees for a 6mm mandrel).
  • the braid pattern was selected and the monofilaments were braided off the spool onto the mandrel by the braiding machine. Tiewraps were used on the end of each mandrel to keep the tension on the filaments, which can be important for heat annealing and obtaining high modulus properties.
  • the braided polymer was heat annealed on the mandrel, and then cut into desired lengths with a blade and removed from the mandrel.
  • a representative implant was measured to have an outer diameter of about 6.0mm and a length of about 20mm.
  • Implants were coated with a support coating made from poly(glycolide-co- capro lactone) (PGCL) cured with hexamethylene diisocyanate.
  • PGCL 50:50 copolymer was prepared as follows. A 100 mL round-bottom flask was dried in oven at 1 10 °C and then cooled to room temperature under a nitrogen atmosphere. The flask was charged with Sn(Oci (15 mg), pentaerythritol (68 mg), glycolide (10.0 g), and ⁇ -caprolactone (10.0 g), respectively. Subsequently, the flask was equipped with a magnetic stir bar and a three-way valve connected to a nitrogen balloon.
  • the four-arm PGCL 50:50 (1.0 g) and HDI (375 ⁇ ) were dissolved in 20 mL dichloromethane to make a stock solution for spray-coating.
  • a steel mandrel of 2 mm in diameter was mounted vertically onto a mechanical stirrer, and the braided implant was placed over the mandrel.
  • a spray gun (Badger 150) was arranged perpendicular to the mandrel, and connected to a nitrogen cylinder and a reservoir containing the stock solution. The mechanical stirrer was turned on to spin the mandrel and the solution was sprayed onto the braid by applying the nitrogen flow. The coating weight could be controlled by the spray time.
  • a catalyst such as tin octanoate, zinc octanoate, aluminum tris(acetylacetonate), etc. may also be used in the curing process to reduce the curing time and/or temperature.
  • Coated devices were placed in an MSI radial force tester to obtain RRF and COF values, both measured in mm Hg, at time points prior to and subsequent to crimping the device to 2.5mm.
  • Implant diameter recovery was also affected by the percent gel content in the coating, as shown in Fig. 8.
  • Gel content is defined as the quantity of coating that is sufficiently crosslinked so that it is no longer soluble in a solvent.
  • the calculation of gel content in the PGCL/HDI coating is described in Example 3. The inventors have found in this Example that a coating gel content of greater than about 80% is preferred to achieve a diameter recovery of at least about 90%.
  • Table I shows the percent recovery, radial resistive force (RRF) at a post-crimp diameter of 4.5mm, and chronic outward force (COF) at a post-crimp diameter of 4.5mm, for coated implants with varying coating weight.
  • RRF radial resistive force
  • COF chronic outward force
  • the gel content of a coated implant (such as the implant described in Example 1) was measured via extraction.
  • the PGCL/HDI coated device was placed in 5mL of
  • % gel content in coating ((mass of coated device after extraction - mass of uncoated device) / (mass of coated device before extraction - mass of uncoated device)) x 100.
  • a control experiment on the uncoated woven PLGA 10:90 structure showed no appreciable mass loss in a similar experiment.
  • a pattern was laser cut into a tubular base material to produce self-expanding medical implants. Some of the implants were coated with a support coating made from PGCL cured with HDI, as described in Example 1. Similar to the previous examples, coated implants demonstrated higher RRF and COF properties as compared with uncoated implants.
  • Braided implants having an as-manufactured diameter of 6mm were manufactured using a PLGA 75:25 copolymer using a manufacturing process similar to that of Example 1.
  • the implants were coated with a support coating made from PLCL 50:50 prepared as follows.
  • a 250mL round-bottom flask was dried in an oven at 110°C and cooled to room temperature in a nitrogen atmosphere.
  • the flask was charged with Sn(Oci (1 1.5 mg), pentaerythritol (204 mg), lactide (30.0 g), and ⁇ -caprolactone (30.0 g), respectively. Subsequently, the flask was equipped with a magnetic stir bar and a three-way valve connected to a nitrogen balloon.
  • Coated devices were crimped to 1.85mm in an MSI radial force tester (Model# RX550-100) to obtain RRF and COF values, both measured in mm Hg, at a diameter of 4.5mm for 6 mm device and at a diameter of 5.5mm for 7mm device.
  • RRF and COF were found to be directly proportional to the coating weight.
  • the inventors note that both RRF and COF for the coated devices are significantly higher than for uncoated devices.
  • a 7 mm implant having an increase in weight due to the PLCL/HDI coating of about 45% resulted in RRF of about 450mm Hg and COF of 90 mm Hg, measured at 4.5 mm, while the uncoated device resulted in RRF and COF of 80mm and 0mm Hg, respectively.
  • Braided implants having an as-manufactured diameter of 6mm were manufactured using a PLGA 75:25 copolymer using a manufacturing process similar to that of Example 1.
  • the implants were coated with a support coating made from poly trimethylene carbonate (PTMC) and hexamethylenediisocyante.
  • PTMC poly trimethylene carbonate
  • the PTMC three arm polymer was prepared as follows. A 100 mL round-bottom flask, previously dried under heat and vacuum, was charged with Sn(Oct)2 (20 mg), triethanolamine(298.4 mg) and trimethylene carbonate (30g) respectively. Subsequently, the flask was equipped with a magnetic stir bar and a three-way valve connected to a nitrogen balloon.
  • the flask was thoroughly degassed under reduced pressure and flushed with nitrogen and then placed into an oil bath which was preheated to 70°C. The oil bath temperature was then increased to 100°C over 15 minutes. The reaction was stirred at 100°C for 23 h under a nitrogen atmosphere. After cooling to room temperature, the viscous liquid obtained was dissolved overnight in approximately 50mL dichloromethane and subsequently precipitated from 550 mL ethanol. The precipitated polymer was stirred for one hour after which the ethanol was decanted. The process of dissolving the polymer in dichloromethane and precipitating in ethanol was repeated.
  • the polymer was then dissolved in dichloromethane, precipitated into 550 mL diethyl ether and stirred for one hour after which time the diethyl ether was decanted. The polymer was then dried under vacuum at 70 °C for a period of 72 hours. Typically 24g of polymer was recovered using above process. GPC characterization of the final polymer revealed a number average molecular weight (Mn) of 29kDa and a polydispersity index (PDI) of 2.0.
  • Mn number average molecular weight
  • PDI polydispersity index
  • RRF radial resistive force
  • COF chronic outward force
  • PTMC/HDI coated devices a 6 mm implant having an increase in weight due to the PTMC/HDI support coating of about 50% resulted in RRF of 490mm Hg and COF of 83 mm Hg, measured at 4.5 mm.
  • PTMC is considered to be a surface-degrading material which does not generate acidic by-products upon degradation.
  • the support coating solution was sprayed onto PLGA 85: 15 braided implants. After spray coating, the implants were allowed to dry for up to 60 minutes in a nitrogen atmosphere. The implants were then transferred to a curing oven for a two-step curing process consisting of a first step at 75°C for 16 hours, followed by a second step at 100°C for a minimum of 96 hours.
  • the PLGA stent structure had an initial diameter of 7 mm, which was crimped to a diameter of 1.9 mm using an MSI radial force tester and then allowed to recover its original diameter.
  • the radial forces of the devices at 5.5 mm were recorded.
  • both RRF and COF decrease with the increase of DDI content in the elastomers, indicating that the tensile stress of the coating elastomers decreases in the same order.
  • Coated devices were crimped to 1.85mm in an MSI radial force tester (Model# RX550-100) to obtain RRF and COF values, both measured in mm Hg, at a diameter of 5.5mm.
  • the addition of more DDI results in looser elastomer network (i.e. lower crosslink density). As a consequence, softer elastomers with lower tensile stress were obtained.
  • both tensile stress and break strain of these elastomers may be tailored by using chain terminators such as monoisocyanates.
  • thermoset elastomers The films were cured at 100 °C for 72 hours to yield corresponding thermoset elastomers.
  • the fully cured films were subject to tensile-strain test on an INSTRON instrument.
  • swelling index (SI) was measured by immersing the films in chloroform at room temperature for 1 hour and calculating SI using the following equation.
  • WQ and Ws are the mass of films before and after swelling, respectively, and po and PCHCB are the density of the elastomers and chloroform, respectively.
  • po and PCHCB are the density of the elastomers and chloroform, respectively.
  • DD may be used as a chain terminator to reduce the crosslink density of thermoset elastomers. From the data shown in Table II, we can see that the elongation at break of the films used in this example increases 25% when DD is added to the elastomer. Meanwhile, the maximum tensile stress of films containing DD decreases 20% as compared with their DD-free analogs. In addition, SI, which is a measurement of crosslink density, indicates a lower crosslink density for films containing DD. These results demonstrate that DD can play an important role in adjusting the crosslink density and mechanical properties of the elastomers.
  • DD may also change the wettability of elastomers.
  • the hydrophobicity of the elastomers increases considerably when hydrophobic DD molecules are engrafted onto the elastomers, as evidenced by the water contact angle change shown in Table II. Consequently, the water penetration and elastomer degradation would be attenuated by hydrophobicity increase.
  • 1-dodecanol (DD) was used to manipulate the chemical crosslink density and consequently the mechanical properties of support coatings made from 4-arm PLCL and HDI- based elastomers.
  • Dichloromethane solutions of PLCL prepolymer and HDI (10: 1 wt/wt PLCL:HDI) and DD (1 :56 wt/wt DD /PLCL) were spray coated onto a braided PLGA 85: 15 stent structure (diameter 7 mm) and cured in-situ.
  • thermoset elastomer support coatings of the present invention resulted in the ability to maintain elastomer strength over prolonged periods of time.
  • Stents were deployed during bench testing into tubing designed to simulate the size and compliance of a superficial femoral artery, and the stents and tubing were exposed to pH neutral saline at 37°C.
  • stents were removed from the tubing, washed in deionized water, dried and tested.
  • RRF and COF values were determined using an MSI radial force tester, which compressed the stents to 5mm and then allowed for re-expansion to nominal deployment diameters.
  • the PLCL material described in this example with DD as a chain terminator applied to stent structures described in Example 11 resulted in the ability to maintain at least 75% of RRF after four weeks and 50% of RRF after twelve weeks.
  • stents coated with the PLCL material without DD as a chain terminator were able to maintain only about 50% of RRF after four weeks and about 10% of RRF after twelve weeks.
  • the PLCL material described in this example with DD as a chain terminator applied to stent structures described in Example 1 1 did not show a decrease in COF through eight weeks, and maintained at least about 50% of COF through twelve weeks.
  • the stents coated with the PLCL material without DD as a chain terminator were able to maintain only about 20% of COF through eight weeks, and had a negligible COF at twelve weeks.
  • the devices described in this Example were generally constrained in a reduced configuration of about 55% to about 90% of their as-manufactured diameters while being held at 37°C for the described time periods.
  • the properties described in this Example are representative of devices having as -manufactured diameters of about 7mm, that are held at reduced diameters of about 4mm to about 6.5mm over the described time periods.
  • An additional means to assess scaffold performance is to measure its resistance to point compressive loads, which is obtained by generating a point load across the longitudinal axis of the scaffold at 50% compression of the scaffold outer diameter. Resistance to point compression is measured in Newtons (N). Testing of devices manufactured from 0.005" and 0.006" fibers resulted in the data shown in Table III, below:
  • Table III Resistance to point compression of devices of the present invention manufactured from 0.005" and 0.006" diameter fibers
  • Braided tubular implants were prepared in accordance Example 1.
  • the implants each comprised 32 strands of 75:25 poly(lactic acid co-glycolic acid) copolymer (PLGA) having a weight percentage of about 75% lactic acid, braided into 7mm diameter tubes and cut into 20mm lengths.
  • the as -manufactured braided implants were coated with a conformal support coating comprising 50:50 poly(lactic acid-co-caprolactone) copolymer (PLCL) crosslinked with hexamethylene diisocyanate.
  • the coated implants were further coated with an additional conformal coating comprising a mixture of PLGA and paclitaxel as active agent.
  • the PLGA in the paclitaxel-containing coating comprised about 50-85% lactic acid, with the balance being glycolic acid.
  • the amount of paclitaxel carried by each implant was controlled by the thickness and loading rate of the paclitaxel-containing coating.
  • thickness By controlling thickness to between ⁇ 1 ⁇ to 8 ⁇ and loading rate from about 0.3wt% to about 5wt%, the inventors have found a drug loading for a 7mm diameter implant to preferably be about 1.5-25Mg per 10mm of implant length, more preferably 2-20 ⁇ g per 10mm of implant length, more preferably 2.5-15Mg per 10mm of implant length, and yet more preferably 5- 13Mg per 10mm of implant length. The inventors believe that such loading rates yield preferred drug release profiles.
  • a drug release duration of about 30-120 days is preferred, about 45-90 days more preferred, and about 45-50 days is yet more preferred, to be used in vascular applications to inhibit neointimal formation and to be fully released prior to significant device degradation and associated loss of integrity.
  • Implants manufactured in accordance with this Example were placed in phosphate buffered saline (PBS) buffer with sodium dodecyl sulfate (SDS) surfactant to measure the release of paclitaxel over time. The results are shown in Fig. 14. Implants 1, 2 and 3 were loaded with 32 ⁇ , 2 ⁇ g, and 12 ⁇ of paclitaxel, respectively. Implants 2 and 3 fully released all paclitaxel within about 75 days and 150 days, respectively. In contrast, implant 1 released only about 12 ⁇ of paclitaxel (or about 38% of its loading) after 75 days.
  • PBS phosphate buffered saline
  • SDS sodium dodecyl sulfate
  • Implants of the present invention were compared to several commercially available, biostable self-expanding implants, namely the VIABAHN ® Endoprosthesis (a self-expanding nitinol stent-graft from W.L. Gore & Associates, Inc.), the S.M.A.R.T. ® stent (a self-expanding nitinol stent from Cordis Corp.), and the WALLSTENT ® Endoprosthesis (a self-expanding steel stent from Boston Scientific Corp.).
  • VIABAHN ® Endoprosthesis a self-expanding nitinol stent-graft from W.L. Gore & Associates, Inc.
  • S.M.A.R.T. ® stent a self-expanding nitinol stent from Cordis Corp.
  • WALLSTENT ® Endoprosthesis a self-expanding steel stent
  • the RRF and COF values of these commercially available implants were measured and compared with 6mm (having a 127° braid angle) and 7mm (having braid angles of 127° and 110°) diameter implants made from PLGA 75:25.
  • the PLGA implants were made using a manufacturing process similar to that of Example 1, and coated with a support coating made from PLCL 50:50 copolymer and hexamethylenediisocyanate as described in Example 5 with a coating weight of between 44% and 50% for the 7mm devices and a weight of 47% to 57% for the 6mm devices.
  • RRF and COF values were measured at post-crimp diameters corresponding to the range of intended target vessel diameters. The results demonstrated that the implants of the present invention are characterized by mechanical properties such as RRF and COF that are comparable to their commercially available metallic counterparts.
  • Braided implants having an as-manufactured diameter of 6mm were manufactured using a PLGA 75:25 copolymer using a manufacturing process similar to that of Example 1.
  • One set of implants were coated with a support coating comprising PGCL 50:50 and hexamethylenediisocyanate, and another set of implants were coated with a support coating comprising PLCL 50:50 and hexamethylenediisocyanate.
  • the implants were deployed into vessel-compliant tubing and maintained at 37°C under simulated flow conditions for 0, 7, 14, and 28 days. At those time points, the implants were explanted from the tubing, dried overnight, and the COF of the implants was measured at a post-crimp diameter of 4.5mm. The results are shown in Fig.
  • the present invention provides woven and non-woven self-expanding medical implants for placement within a bodily lumen that have sufficient strength and other mechanical properties that are necessary to effectively treat a variety of medical conditions. While aspects of the invention have been described with reference to example embodiments thereof, it will be understood by those skilled in the art that various changes in form and details may be made therein without departing from the scope of the invention.

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Abstract

L'invention concerne des implants médicaux auto-expansibles à placer à l'intérieur d'une lumière d'un patient. Les implants comprennent une structure tissée ou non tissée, ayant une configuration sensiblement tubulaire, et sont conçus pour être à profil bas, de telle sorte qu'ils peuvent être posés avec un cathéter de petit diamètre. Les implants ont une capacité de récupération élevée et des caractéristiques mécaniques souhaitées.
PCT/US2013/036743 2012-04-16 2013-04-16 Implant médical d'élution de médicament WO2013158619A2 (fr)

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