US20070278720A1 - Implantable medical devices made from polymer-bioceramic composite - Google Patents

Implantable medical devices made from polymer-bioceramic composite Download PDF

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US20070278720A1
US20070278720A1 US11/725,630 US72563007A US2007278720A1 US 20070278720 A1 US20070278720 A1 US 20070278720A1 US 72563007 A US72563007 A US 72563007A US 2007278720 A1 US2007278720 A1 US 2007278720A1
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polymer
particles
composite
bioceramic particles
bioceramic
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US11/725,630
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Yunbing Wang
David Gale
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Abbott Cardiovascular Systems Inc
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Abbott Cardiovascular Systems Inc
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Priority claimed from US11/523,866 external-priority patent/US20070282434A1/en
Priority claimed from US11/529,996 external-priority patent/US7842737B2/en
Application filed by Abbott Cardiovascular Systems Inc filed Critical Abbott Cardiovascular Systems Inc
Priority to US11/725,630 priority Critical patent/US20070278720A1/en
Assigned to ABBOTT CARDIOVASCULAR SYSTEMS INC. reassignment ABBOTT CARDIOVASCULAR SYSTEMS INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: WANG, YUNBING, GALE, DAVID C.
Priority to PCT/US2007/017237 priority patent/WO2008016670A2/en
Publication of US20070278720A1 publication Critical patent/US20070278720A1/en
Abandoned legal-status Critical Current

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    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61FFILTERS 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
    • A61F2/00Filters 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
    • A61F2/82Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61LMETHODS 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
    • A61L31/00Materials 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
    • A61L31/12Composite materials, i.e. containing one material dispersed in a matrix of the same or different material
    • A61L31/125Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix
    • A61L31/127Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix containing fillers of phosphorus-containing inorganic materials
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61LMETHODS 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
    • A61L31/00Materials 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
    • A61L31/12Composite materials, i.e. containing one material dispersed in a matrix of the same or different material
    • A61L31/125Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix
    • A61L31/128Composite materials, i.e. containing one material dispersed in a matrix of the same or different material having a macromolecular matrix containing other specific inorganic fillers not covered by A61L31/126 or A61L31/127
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61LMETHODS 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
    • A61L31/00Materials 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
    • A61L31/14Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
    • A61L31/148Materials at least partially resorbable by the body
    • BPERFORMING OPERATIONS; TRANSPORTING
    • B29WORKING OF PLASTICS; WORKING OF SUBSTANCES IN A PLASTIC STATE IN GENERAL
    • B29CSHAPING OR JOINING OF PLASTICS; SHAPING OF MATERIAL IN A PLASTIC STATE, NOT OTHERWISE PROVIDED FOR; AFTER-TREATMENT OF THE SHAPED PRODUCTS, e.g. REPAIRING
    • B29C48/00Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor
    • B29C48/022Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor characterised by the choice of material
    • B29C48/023Extruding materials comprising incompatible ingredients
    • BPERFORMING OPERATIONS; TRANSPORTING
    • B29WORKING OF PLASTICS; WORKING OF SUBSTANCES IN A PLASTIC STATE IN GENERAL
    • B29CSHAPING OR JOINING OF PLASTICS; SHAPING OF MATERIAL IN A PLASTIC STATE, NOT OTHERWISE PROVIDED FOR; AFTER-TREATMENT OF THE SHAPED PRODUCTS, e.g. REPAIRING
    • B29C48/00Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor
    • B29C48/03Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor characterised by the shape of the extruded material at extrusion
    • B29C48/05Filamentary, e.g. strands
    • BPERFORMING OPERATIONS; TRANSPORTING
    • B29WORKING OF PLASTICS; WORKING OF SUBSTANCES IN A PLASTIC STATE IN GENERAL
    • B29CSHAPING OR JOINING OF PLASTICS; SHAPING OF MATERIAL IN A PLASTIC STATE, NOT OTHERWISE PROVIDED FOR; AFTER-TREATMENT OF THE SHAPED PRODUCTS, e.g. REPAIRING
    • B29C48/00Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor
    • B29C48/03Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor characterised by the shape of the extruded material at extrusion
    • B29C48/07Flat, e.g. panels
    • B29C48/08Flat, e.g. panels flexible, e.g. films
    • BPERFORMING OPERATIONS; TRANSPORTING
    • B29WORKING OF PLASTICS; WORKING OF SUBSTANCES IN A PLASTIC STATE IN GENERAL
    • B29CSHAPING OR JOINING OF PLASTICS; SHAPING OF MATERIAL IN A PLASTIC STATE, NOT OTHERWISE PROVIDED FOR; AFTER-TREATMENT OF THE SHAPED PRODUCTS, e.g. REPAIRING
    • B29C48/00Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor
    • B29C48/03Extrusion moulding, i.e. expressing the moulding material through a die or nozzle which imparts the desired form; Apparatus therefor characterised by the shape of the extruded material at extrusion
    • B29C48/09Articles with cross-sections having partially or fully enclosed cavities, e.g. pipes or channels
    • YGENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10STECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10S977/00Nanotechnology
    • Y10S977/70Nanostructure
    • Y10S977/734Fullerenes, i.e. graphene-based structures, such as nanohorns, nanococoons, nanoscrolls or fullerene-like structures, e.g. WS2 or MoS2 chalcogenide nanotubes, planar C3N4, etc.
    • Y10S977/753Fullerenes, i.e. graphene-based structures, such as nanohorns, nanococoons, nanoscrolls or fullerene-like structures, e.g. WS2 or MoS2 chalcogenide nanotubes, planar C3N4, etc. with polymeric or organic binder
    • YGENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10STECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10S977/00Nanotechnology
    • Y10S977/70Nanostructure
    • Y10S977/773Nanoparticle, i.e. structure having three dimensions of 100 nm or less
    • Y10S977/775Nanosized powder or flake, e.g. nanosized catalyst
    • Y10S977/776Ceramic powder or flake
    • YGENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10STECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10S977/00Nanotechnology
    • Y10S977/70Nanostructure
    • Y10S977/831Of specified ceramic or electrically insulating compositions
    • YGENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10TECHNICAL SUBJECTS COVERED BY FORMER USPC
    • Y10STECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y10S977/00Nanotechnology
    • Y10S977/902Specified use of nanostructure
    • Y10S977/904Specified use of nanostructure for medical, immunological, body treatment, or diagnosis
    • Y10S977/931Medical device coating

Definitions

  • This invention relates to implantable medical devices and methods of fabricating implantable medical devices.
  • This invention relates to radially expandable endoprostheses, which are adapted to be implanted in a bodily lumen.
  • An “endoprosthesis” corresponds to an artificial device that is placed inside the body.
  • a “lumen” refers to a cavity of a tubular organ such as a blood vessel.
  • a stent is an example of such an endoprosthesis.
  • Stents are generally cylindrically shaped devices, which function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels.
  • Stepnosis refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce body vessels and prevent restenosis following angioplasty in the vascular system.
  • Restenosis refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.
  • the treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent.
  • Delivery refers to introducing and transporting the stent through a bodily lumen to a region, such as a lesion, in a vessel that requires treatment.
  • Delivery corresponds to the expanding of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into a bodily lumen, advancing the catheter in the bodily lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.
  • the stent In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves compressing or crimping the stent onto the balloon. The stent is then expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn.
  • the stent In the case of a self-expanding stent, the stent may be secured to the catheter via a constraining member such as a retractable sheath or a sock. When the stent is in a desired bodily location, the sheath may be withdrawn which allows the stent to self-expand.
  • the stent must be able to satisfy a number of mechanical requirements.
  • the stent must be capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of a vessel. Therefore, a stent must possess adequate radial strength.
  • Radial strength which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around a circumferential direction of the stent. Radial strength and rigidity, therefore, may also be described as, hoop or circumferential strength and rigidity.
  • the stent Once expanded, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. For example, a radially directed force may tend to cause a stent to recoil inward. Generally, it is desirable to minimize recoil.
  • the stent must possess sufficient flexibility to allow for crimping, expansion, and cyclic loading. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure.
  • the stent must be biocompatible so as not to trigger any adverse vascular responses.
  • the structure of a stent is typically composed of scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts or bar arms.
  • the scaffolding can be formed from wires, tubes, or sheets of material rolled into a cylindrical shape.
  • the scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment).
  • a conventional stent is allowed to expand and contract through movement of individual structural elements of a pattern with respect to each other.
  • a medicated stent may be fabricated by coating the surface of either a metallic or polymeric scaffolding with a polymeric carrier that includes an active or bioactive agent or drug.
  • Polymeric scaffolding may also serve as a carrier of an active agent or drug.
  • a stent may be biodegradable.
  • the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Therefore, stents fabricated from biodegradable, bioabsorbable, and/or bioerodable materials such as bioabsorbable polymers should be configured to completely erode only after the clinical need for them has ended.
  • a potential problem with polymeric stents is that their struts or bar arms can crack during crimping and expansion. This is especially the case with brittle polymers.
  • the localized portions of the stent pattern subjected to substantial deformation during crimping and expansion tend to be the most vulnerable to failure.
  • a stent it is desirable for a stent to have flexibility and resistance to cracking during deployment. It is also advantageous for a stent to be rigid and resistant to creep after deployment. It would also be desirable to be able to control the degradation rate of the device.
  • Certain embodiments of the present invention include a method of fabricating an implantable medical device comprising: treating the surface of a plurality of bioceramic particles with a non-reactive surface modifier; forming a composite including the treated bioceramic particles dispersed within a polymer, the non-reactive surface modifier reducing the agglomeration of the treated bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • Additional embodiments of the present invention include a stent comprising: a structural element including a bioceramic/polymer composite, the composite having plurality of bioceramic particles dispersed within a polymer, a non-reactive surface modifier being on a surface of at least some of the particles, wherein the surface modifier reduces agglomeration of the bioceramic particles during the formation of the composite.
  • Certain other embodiments of the present invention include a method of fabricating an implantable medical device comprising: subjecting a mixture of bioceramic particles and polymer to high sheer stress using twin-screw extruder or a batch mixer; forming a composite including the bioceramic particles dispersed within a polymer, the mechanical smash, reducing the agglomeration of the bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • Some embodiments of the present invention include a method of fabricating an implantable medical device comprising: treating the surface of a plurality of bioceramic particles with a non-reactive surface modifier; subjecting a mixture of the treated bioceramic particles and polymer to high sheer stress using twin-screw extruder or a batch mixer; forming a composite including the treated bioceramic particles dispersed within a polymer, the non-reactive surface modifier or mechanical smash reducing the agglomeration of the treated bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • FIG. 1 depicts a three-dimensional view of a stent.
  • FIG. 2A depicts a section of a structural element from the stent depicted in FIG. 1 .
  • FIG. 2B depicts bioceramic particles dispersed within a polymer matrix.
  • FIG. 3 depicts a schematic plot of the crystal nucleation rate, the crystal growth rate, and the overall rate of crystallization for a semicrystalline polymer.
  • FIG. 4 is a graph depicting the peak stress for PLLA and PLLA/HAP composites.
  • FIG. 5 is a graph depicting the break strain for PLLA and PLLA/HAP composites.
  • FIG. 6 is a graph depicting Young's modulus for PLLA and PLLA/HAP composite stent in compression testing.
  • FIG. 7 is a graph depicting the radial stress for PLGA and PLGA/HAP composite stent in compression testing.
  • FIG. 8 is a graph depicting the compression modulus for PLGA and PLGA/HAP composite stent in compression testing.
  • FIG. 9 is a graph depicting recoil testing of PLGA and PLGA/HAP composite stent.
  • FIGS. 10 A-D depict photographic images of a PLGA/calcium sulfate composite stent with a weight ratio of PLGA/calcium sulfate of 100:1 at zero time point.
  • FIGS. 11 A-D depict photographic images of a PLGA/calcium sulfate composite stent with a weight ratio of PLGA/calcium sulfate of 100:1 at 16 hours of accelerated aging.
  • FIG. 12A depicts a SEM photographic image of LPLG/unmodified calcium sulfate composite before surface modification
  • FIG. 12B depicts a SEM photographic image of LPLG/PEO-b-PPO-b-PEO modified calcium sulfate composite after surface modification
  • FIG. 12C depicts a SEM photographic image of LPLG/stearic acid modified calcium sulfate composite after surface modification
  • FIG. 13A depicts a photographic image of LPLG/calcium sulfate composite formed using twin screw extruder
  • FIG. 13B depicts a photographic image of LPLG/calcium sulfate composite formed using single screw extruder
  • FIG. 13C depicts a photographic image of LPLG/PEO-b-PPO-b-PEO modified calcium sulfate composite formed using twin screw extruder
  • the “glass transition temperature,” Tg is the temperature at which the amorphous domains of a polymer change from a brittle vitreous state to a solid deformable or ductile state at atmospheric pressure.
  • the Tg corresponds to the temperature where the onset of segmental motion in the chains of the polymer occurs.
  • Tg of a given polymer can be dependent on the heating rate and can be influenced by the thermal history of the polymer. Furthermore, the chemical structure of the polymer heavily influences the glass transition by affecting mobility.
  • Stress refers to force per unit area, as in the force acting through a small area within a plane. Stress can be divided into components, normal and parallel to the plane, called normal stress and shear stress, respectively. True stress denotes the stress where force and area are measured at the same time. Conventional stress, as applied to tension and compression tests, is force divided by the original gauge length.
  • “Strength” refers to the maximum stress along an axis which a material will withstand prior to fracture. The ultimate strength is calculated from the maximum load applied during the test divided by the original cross-sectional area.
  • Modulus may be defined as the ratio of a component of stress or force per unit area applied to a material divided by the strain along an axis of applied force that results from the applied force.
  • a material has both a tensile and a compressive modulus.
  • a material with a relatively high modulus tends to be stiff or rigid.
  • a material with a relatively low modulus tends to be flexible.
  • the modulus of a material depends on the molecular composition and structure, temperature of the material, amount of deformation, and the strain rate or rate of deformation. For example, below its Tg, a polymer tends to be brittle with a high modulus. As the temperature of a polymer is increased from below to above its Tg, its modulus decreases.
  • Stress refers to the amount of elongation or compression that occurs in a material at a given stress or load.
  • Elongation may be defined as the increase in length in a material which occurs when subjected to stress. It is typically expressed as a percentage of the original length.
  • “Toughness” is the amount of energy absorbed prior to fracture, or equivalently, the amount of work required to fracture a material.
  • One measure of toughness is the area under a stress-strain curve from zero strain to the strain at fracture. Thus, a brittle material tends to have a relatively low toughness.
  • solvent is defined as a substance capable of dissolving or dispersing one or more other substances or capable of at least partially dissolving or dispersing the substance(s) to form a uniformly dispersed solution at the molecular- or ionic-size level.
  • the solvent should be capable of dissolving at least 0.1 mg of the polymer in 1 ml of the solvent, and more narrowly 0.5 mg in 1 ml at ambient temperature and ambient pressure.
  • an “implantable medical device” includes, but is not limited to, self-expandable stents, balloon-expandable stents, stent-grafts, implantable cardiac pacemakers and defibrillators; leads and electrodes for the preceding; implantable organ stimulators such as nerve, bladder, sphincter and diaphragm stimulators, cochlear implants; prostheses, vascular grafts, grafts, artificial heart valves and cerebrospinal fluid shunts.
  • An implantable medical device can be designed for the localized delivery of a therapeutic agent.
  • a medicated implantable medical device may be constructed by coating the device with a coating material containing a therapeutic agent.
  • the substrate of the device may also contain a therapeutic agent.
  • FIG. 1 depicts a three-dimensional view of stent 100 .
  • a stent may include a pattern or network of interconnecting structural elements 110 .
  • Stent 100 may be formed from a tube (not shown).
  • Stent 100 includes a pattern of structural elements 110 , which can take on a variety of patterns.
  • the structural pattern of the device can be of virtually any design.
  • the embodiments disclosed herein are not limited to stents or to the stent pattern illustrated in FIG. 1 .
  • the embodiments are easily applicable to other patterns and other devices.
  • the variations in the structure of patterns are virtually unlimited.
  • a stent such as stent 100 may be fabricated from a tube by forming a pattern with a technique such as laser cutting or chemical etching.
  • a pattern may include portions of structural elements or struts that are straight or relatively straight, an example being a portion 120 .
  • patterns may include structural elements or struts that include curved or bent portions such as portions 130 , 140 , and 150 .
  • An implantable medical device can also be made partially or completely from a biodegradable, bioabsorbable, or biostable polymer.
  • a polymer for use in fabricating an implantable medical device can be biostable, bioabsorbable, biodegradable or bioerodable.
  • Biostable refers to polymers that are not biodegradable.
  • the terms biodegradable, bioabsorbable, and bioerodable are used interchangeably and refer to polymers that are capable of being completely degraded and/or eroded when exposed to bodily fluids such as blood and can be gradually resorbed, absorbed, and/or eliminated by the body.
  • the processes of breaking down and absorption of the polymer can be caused by, for example, hydrolysis and metabolic processes.
  • polymers tend to have a number of shortcomings for use as materials for implantable medical devices such as stents.
  • Many biodegradable polymers have a relatively low modulus at the physiological conditions in the human body.
  • the strength to weight ratio of polymers is smaller than that of metals.
  • a polymeric stent with inadequate radial strength can result in mechanical failure or recoil inward after implantation into a vessel.
  • a polymeric stent requires significantly thicker struts than a metallic stent, which results in an undesirably large profile.
  • polymers such as biodegradable polymers
  • Tg above human body temperature which is approximately 37° C.
  • These polymer systems exhibit a brittle fracture mechanism in which there is little or no plastic deformation prior to failure.
  • a stent fabricated from such polymers can have insufficient toughness for the range of use of a stent.
  • Creep refers to the gradual deformation that occurs in a polymeric contruct subjected to an applied load. Creep occurs even when the applied load is constant.
  • Stress relaxation is also a consequence of delayed molecular motions as in creep. Contrary to creep, however, which is experienced when the load is constant, stress relaxation occurs when deformation (or strain) is constant and is manifested by a reduction in the force (stress) required to maintain a constant deformation.
  • Physical aging refers to densification in the amorphous regions of a semi-crystalline polymer. Densification is the increase in density of a material or region of a material. Densification, and thus physical aging, is also the result of relaxation or rearrangement of polymer chains.
  • Various embodiments of the present invention include an implantable medical device fabricated from a composite including a polymer matrix or continuous phase and bioceramic particles as a discrete phase.
  • the bioceramic particles may tend to reduce or eliminate a number of the above-mentioned shortcomings of polymers.
  • the bioceramic particles can increase the toughness and modulus and modify the degradation rate of the polymer.
  • the composite may include a plurality of bioceramic particles dispersed within the polymer.
  • the bioceramic particles are uniformly dispersed throughout the biodegradable polymer.
  • a uniform dispersion can result in a uniform increase in toughness and modulus and modification of degradation rate.
  • the bioceramic particles are uniformly or substantially uniformly dispersed within the biodegradable polymer.
  • a structural element of an implantable medical device may be fabricated from a bioceramic/polymer composite.
  • Structural elements can include, but are not limited to, any supporting element such as a strut, wire, or filament.
  • FIG. 2A depicts a section 200 of a structural element 110 from stent 100 . A portion 210 of section 200 is shown in an expanded view in FIG. 2B .
  • FIG. 2B depicts bioceramic particles 220 dispersed throughout a polymer matrix 230 .
  • Bioceramics can include any ceramic material that is compatible with the human body. More generally, bioceramic materials can include any type of compatible inorganic material or inorganic/organic hybrid material. Bioceramic materials can include, but are not limited to, alumina, zirconia, apatites, calcium phosphates, silica based glasses, or glass ceramics, and pyrolytic carbons. Bioceramic materials can be bioabsorbable and/or active. A bioceramic is active if it actively takes part in physiological processes. A bioceramic material can also be “inert,” meaning that the material does not absorb or degrade under physiological conditions of the human body and does not actively take part in physiological processes.
  • apatites and other calcium phosphates include, but are not limited hydroxyapatite (Ca 10 (PO 4 ) 6 (OH) 2 ), floroapatite (Ca 10 (PO 4 ) 6 F 2 ), carbonate apatide (Ca 10 (PO 4 ) 6 CO 3 ), tricalcium phosphate (Ca 3 (PO 4 ) 2 ), octacalcium phosphate (Ca 8 H 2 (PO 4 )6-5H 2 O), octacalcium phosphate (Ca 8 H 2 (PO 4 )6-5H 2 O), calcium pyrophosphate (Ca 2 P 2 O 7 -2H 2 O), tetracalcium phosphate (Ca 4 P 2 O 9 ), and dicalcium phosphate dehydrate (CaHPO 4 -2H 2 O).
  • bioceramics can also include bioactive glasses that are bioactive glass ceramics composed of compounds such as SiO 2 , Na 2 O, CaO, and P 2 O 5 .
  • bioactive glasses that are bioactive glass ceramics composed of compounds such as SiO 2 , Na 2 O, CaO, and P 2 O 5 .
  • Bioglass® a commercially available bioactive glass, Bioglass®, is derived from certain compositions of SiO 2 —Na2O—K 2 O—CaO—MgO—P 2 O 5 systems.
  • Some commercially available bioactive glasses include, but are not limited to:
  • Another commercially available glass ceramic is A/W.
  • bioceramic particles in a composite implantable medical device may be used to inhibit or prevent infection since some bioceramics can have an anti-infective property.
  • Bioceramics may release various ions such as calcium and phosphate ions which broadly exist in human body fluid and blood plasma. Examples of bioceramics that release calcium and/or phosphate ions include various calcium phosphates and bioactive glasses. The released ions may depress foreign body reaction. Trends Biomater. Artif. Tren, Vol 18 (1), pp 9-17.
  • an implantable medical device such as a stent can be medicated by incorporating an active agent in a coating over the device or within the substrate of the device.
  • the ions released from bioceramics can have an additive therapeutic and/or a synergistic therapeutic effect to the active agent.
  • ions can be used in conjunction with anti-proliferative and/or anti-inflammatory agents.
  • Bioceramic particles can be partially or completely made from a biodegradable, bioabsorbable, or biostable ceramic.
  • bioabsorbable bioceramics include various types of bioglass materials, tetracalcium phosphate, amorphous calcium phosphate, alpha-tricalcium phosphate, and beta-tricalcium phosphate.
  • Biostable bioceramics include alumina and zirconia.
  • the bioceramic particles can include, but are not limited to, nanoparticles and/or micro particles.
  • a nanoparticle refers to a particle with a characteristic length (e.g., diameter) in the range of about 1 nm to about 1,000 nm.
  • a micro particle refers to a particle with a characteristic length in the range of greater than 1,000 nm and less than about 10 micrometers.
  • bioceramic particles can be of various shapes, including but not limited to, spheres and fibers.
  • the particles size distribution can be important in modifying the properties of the polymer. Generally, a narrow size distribution is preferable.
  • the composite of a structural element of a device may have between 0.01% and 10% of bioceramic particles by weight, or more narrowly, between 0.5% and 2% bioceramic particles by weight as compared to the polymer matrix of the composite.
  • bioceramic particles can reduce or eliminate a number of shortcomings of polymers that are used for implantable medical devices.
  • bioceramic particles can increase the fracture toughness of polymers of implantable medical device. In general, the higher the fracture toughness, the more resistant a material is to the propagation of cracks.
  • bioceramic particles may be used in a composite having a matrix polymer that is brittle at physiological conditions. In particular, such a polymer can have a Tg above body temperature. In one embodiment, the bioceramic particles may be nanoparticles.
  • Certain regions of an implantable medical device experience a high degree of stress and strain when the device is under stress during use.
  • curved or bending regions such as portions 130 , 140 , and 150 can have highly concentrated strain which can lead to fracture.
  • the bioceramic particles can increase fracture toughness by reducing the concentration of strain by dispersing the strain over a large volume of the material. Particles can absorb energy due to applied stress and disperse energy about a larger volume in the bioceramic/polymer composite.
  • the stress and strain in a device fabricated from a bioceramic composite is divided into many small interactions involving numerous individual particles.
  • the crack breaks up into finer and finer cracks due to interaction with the particles.
  • the particles tend to dissipate the energy of imparted to the device by the applied stress.
  • For a give weight ratio of particles to matrix as the size of the particles decreases the number of particles dispersed throughout the device per unit volume also increases.
  • the number of particles to disperse the energy of applied stress to the device increases. Therefore, it is advantageous to use nanoparticles to increase the toughness of the polymer. It has been shown that the fracture toughness of a polymeric material can be improved by using nanoparticles as a discrete or reinforcing phase in a composite. J. of Applied Polymer Science, 94 (2004) 796-802.
  • Bioceramic particles by providing more crystallites in a network in the bioceramic/polymer composite increase fracture toughness.
  • bioceramic particles can be used to increase the the strength and modulus of the polymer.
  • a polymeric stent requires a high radial strength in order to provide effective scaffolding of a vessel.
  • Many biodegradable polymers have a relatively low strength and modulus as compared to metals.
  • a composite with bioceramic particles with a higher modulus than a matrix polymer may have a higher strength and modulus than the polymer.
  • the higher strength and modulus may allow for the manufacture of a composite stent with much thinner struts than a stent fabricated from the matrix polymer alone.
  • relatively low modulus polymers include, but are not limited to, poly(D,L-lactide-co-glycolide), poly(lactide-co-caprolactone), poly(lactide-co-trimethylene carbonate), poly(glycolide-co-caprolactone), and poly(D,L-lactide). It has been reported that composites with nanoparticles can increase the modulus of a polymer by 1-2 orders of magnitude. Mechanical Properties of Polymers and Composites, Lawrence E. Nielsen and Robert F. Landel, 2 nd ed., p. 384-385 (1993).
  • bioceramic particles in a polymer composite can also reduce or eliminate creep, stress relaxation, and physical aging. It is believed that particles can act as “net point” that reduce or inhibit movement of polymer chains in amorphous regions of a polymer.
  • the crystallinity of a bioceramic/polymer composite that forms an implantable device can be controlled to reduce or eliminate creep, stress relaxation, and physical aging. As indicated above, these phenomena in a polymer are due to rearrangement or relaxation of polymer chains.
  • a structural element of an implantable medical device may include a composite having a plurality of crystalline domains dispersed within an amorphous biodegradable polymeric matrix phase.
  • the crystalline domains may be formed around bioceramic particles.
  • the composite that makes up the structural element may have a relatively low crystallinity.
  • the crystallinity can be less than 50%, 30%, 20%, or less than 10%.
  • the device can be fabricated so that the resulting composite has a relatively large number of crystalline domains that are relatively small.
  • the average crystal size can be less than 10 microns, or less than 5 microns.
  • the polymer may become less brittle and, which increases the fracture toughness.
  • the crystallinity of the resulting polymer can be relatively low, the presence of the relatively large number of relatively small crystalline domains can reduce or eliminate physical aging, creep, and stress relaxation.
  • the size and number of crystallites domains can be controlled during formation of a polymer construct from an implantable medical device is fabricated.
  • Polymer constructs such as tubes, can be formed using various types of forming methods, including, but not limited to extrusion or injection molding.
  • Representative examples of extruders include, but are not limited to, single screw extruders, intermeshing co-rotating and counter-rotating twin-screw extruders, and other multiple screw masticating extruders.
  • a mixture of a polymer and bioceramic particles can be extruded to form a polymer construct, such as a tube.
  • a polymer melt mixed with the bioceramic particles can be conveyed through an extruder and forced through a die in the shape of as an annular film in the shape of a tube.
  • the annular film can be cooled below the melting point, Tm, of the polymer to form an extruded polymeric tube.
  • Tm melting point
  • the annular film may be conveyed through a water bath at a selected temperature.
  • the annular film may be cooled by a gas at a selected temperature.
  • the annular film may be cooled at or near an ambient temperature, e.g. 25° C.
  • the annular film may be cooled at a temperature below ambient temperature.
  • the temperature of the polymer construct during cooling can be between Tg and Tm.
  • the bioceramic particles provide a point of nucleation in the polymer melt for the formation of crystalline domains.
  • a network of many small crystalline domains is formed, which can work to tie crystalline domains together and reduce, inhibit or prevent fracturing, creep, stress relaxation, and physical aging of the polymer.
  • the crystalline domains can serve as net points in the amorphous domains that restrict the freedom of movement of polymer chains in the amorphous domain. As a result, physical aging, creep, and stress relaxation can be reduced. In addition, for the reasons discussed above, the toughness of the polymer is also increased.
  • both microparticles and nanoparticles can be used as nucleation points.
  • the crystalline domains become more effective in increasing fracture toughness and reducing physical aging, creep, and stress relaxation.
  • the size of the crystalline domains can be controlled by the temperature of the cooling polymer construct from an extruder.
  • crystallization tends to occur in a polymer at temperatures between Tg and Tm of the polymer. The rate of crystallization in this range varies with temperature.
  • FIG. 3 depicts a schematic plot of the crystal nucleation rate (R N ), the crystal growth rate (R CG ), and the overall rate of crystallization (R CO ).
  • the crystal nucleation rate is the growth rate of new crystals and the crystal growth rate is the rate of growth of formed crystals.
  • the overall rate of crystallization is the sum of curves R N and R CG .
  • the temperature of the cooling polymer construct can be at a temperature at which the overall crystallization rate is relatively low. At such a temperature, the increase in crystallinity is predominantly due to formation of crystalline domains around the bioceramic particles, rather than the growth of existing crystals.
  • the temperature can be in a range in which the crystal nucleation rate is larger than the crystal growth rate. In one embodiment, the temperature can be in a range in which the crystal nucleation rate is substantially larger than the crystal growth rate. For example, the temperature can be where the ratio of the crystal nucleation rate to crystal growth rate is 2, 5, 10, 50, 100, or greater than 100. In another embodiment, the temperature range may be in range, ⁇ T, shown in FIG. 3 , between about Tg to about 0.25(Tm ⁇ Tg)+Tg.
  • good bonding between a continuous phase and a discrete or reinforcing phase in a composite material facilitates improvement of the mechanical performance of the composite.
  • increase of the modulus and fracture toughness of a polymer due to a bioceramic particle phase can be enhanced by good bonding between the polymer and particles.
  • bioceramic particles may include an adhesion promoter to improve the adhesion between the particles and the polymer matrix.
  • an adhesion promoter can include a coupling agent.
  • a coupling agent refers to a chemical substance capable of reacting with both the bioceramic particle and the polymer matrix of the composite material. A coupling agent acts as an interface between the polymer and the bioceramic particle to form a chemical bridge between the two to enhance adhesion.
  • the adhesion promoter may include, but is not limited to, silane and non-silane coupling agents.
  • the adhesion promoter may include 3-aminopropyltrimethoxysilane, 3-aminopropyltriethoxysilane, aminopropylmethyldiethoxy silane, organotrialkoxysilanes, titanates, zirconates, and organic acid-chromium chloride coordination complexes.
  • 3-aminopropyltrimethoxysilane has been shown to facilitate adhesion between poly(L-lactide) and bioglass. Biomaterials 25 (2004) 2489-2500.
  • the surface of the bioceramic particles may be treated with an adhesion promoter prior to mixing with the polymer matrix.
  • the bioceramic particles can be treated with a solution containing the adhesion promoter. Treating can include, but is not limited to, coating, dipping, or spraying the particles with an adhesion promoter or a solution including the adhesion promoter. The particles can also be treated with a gas containing the adhesion promoter.
  • treatment of the bioceramic particles includes mixing the adhesion promoter with solution of distilled water and a solvent such as ethanol and then adding bioceramic particles. The bioceramic particles can then be separated from the solution, for example, by a centrifuge, and the particles can be dried.
  • the bioceramic particles may then used to form a polymer composite.
  • the adhesion promoter can be added to the particles during formation of the composite.
  • the adhesion promoter can be mixed with a bioceramic/polymer mixture during extrusion.
  • a device may be composed in whole or in part of materials that degrade, erode, or disintegrate through exposure to physiological conditions within the body until the treatment regimen is completed.
  • the device may be configured to disintegrate and disappear from the region of implantation once treatment is completed.
  • the device may disintegrate by one or more mechanisms including, but not limited to, dissolution and chemical breakdown.
  • the duration of a treatment period depends on the bodily disorder that is being treated. For illustrative purposes only, in treatment of coronary heart disease involving use of stents in diseased vessels, the duration can be in a range from about a month to a few years. However, the duration is typically in a range from about six to twelve months. Thus, it is desirable for an implantable medical device, such as a stent, to have a degradation time at or near the duration of treatment. Degradation time refers to the time for an implantable medical device to substantially or completely erode away from an implant site.
  • bodily conditions can include, but are not limited to, all conditions associated with bodily fluids (contact with fluids, flow of fluids) and mechanical forces arising from body tissue in direct and indirect contact with a device.
  • Degradation of polymeric materials principally involves chemical breakdown involving enzymatic and/or hydrolytic cleavage of device material due to exposure to bodily fluids such as blood.
  • Chemical breakdown of biodegradable polymers results in changes of physical and chemical properties of the polymer, for example, following exposure to bodily fluids in a vascular environment. Chemical breakdown may be caused by, for example, hydrolysis and/or metabolic processes. Hydrolysis is a chemical process in which a molecule is cleaved into two parts by the addition of a molecule of water. Consequently, the degree of degradation in the bulk of a polymer is strongly dependent on the diffusivity, and hence the diffusion rate of water in the polymer.
  • Another deficiency of some biodegradable polymers is that the degradation rate is slow and results in a degradation time of a stent outside of the desired range.
  • a preferred degradation is from six to twelve months.
  • Increasing the equilibrium content of moisture in a biodegradable polymer that degrades by hydrolysis can increase the degradation rate of a polymer.
  • Various embodiments of the present invention include increasing the equilibrium moisture content in a polymer of a device to accelerate the degradation rate.
  • bioabsorbable bioceramic particles may be included in a bioceramic/polymer composite device to increase the degradation rate of the polymer and to decrease the degradation time of a device made from the composite.
  • the degradation rate of a bioceramic/polymer composite device can be tuned and/or adjusted to a desired time frame. As the bioceramic particle erodes within the polymeric matrix, the porosity of the matrix increases. The increased porosity increases the diffusion rate of moisture through the polymeric matrix, and thus, the equilibrium moisture content of the polymeric matrix. As a result, the degradation rate of the polymer is increased. The porous structure also increases the transport of degradation products out of the matrix, which also increases the degradation rate of the matrix.
  • the degradation rate and degradation time of the device can be tuned or controlled through variables such as the type of bioceramic material and the size and shape of particles.
  • bioceramic materials can be selected to have a higher degradation rate than the polymer matrix. The faster the degradation rate of the bioceramic material, the faster the porosity of the polymer matrix increases which results in a greater increase in the degradation rate of the polymer matrix. Additionally, the size of the particles influence the time for erosion of the particles. The smaller the particles, the faster the erosion of the particles because of the higher surface area per unit mass of particles.
  • nanoparticles may have a relatively fast erosion rate compared to microparticles.
  • elongated particles such as fibers
  • short fibers may tend to erode faster than longer fibers.
  • Short fibers refer to long fibers than have been cut into short lengths.
  • the short fibers may be made by forming fibers as described above, and cutting them into short lengths.
  • a length of at least a portion of the short fibers is substantially smaller than a diameter of the formed tube.
  • the short fibers may be less than 0.05 mm long.
  • the short fibers may be between 0.05 and 8 mm long, or more narrowly between 0.1 and 0.4 mm long or 0.3 and 0.4 mm long.
  • the size and distribution of pores created by erosion of bioceramic particles can also influence the degradation rate and time of the polymer matrix.
  • Smaller particles, such as nanoparticles create a porous network that exposes a larger volume of polymer matrix to bodily fluid than larger particles, like microparticles.
  • the degradation rate and time of the matrix may be higher when nanoparticles are used rather than microparticles.
  • the particles and the device can be designed to have a selected erosion rates and degradation time.
  • the particles can designed erode away in several minutes, hours, days, or a month upon exposure to bodily fluid.
  • biodegradable polymers degrade by the mechanism of hydrolysis.
  • the rate of the hydrolysis reaction tends to increase as the pH decreases.
  • the degradation products of such polymers as polylactides are acidic, the degradation products have an autocatalytic effect. Therefore, the pH of the degradation products of the bioceramics can also affect the degradation rate of a device. Therefore, bioceramic particles with acidic degradation by-products may further increase the rate of degradation of a matrix polymer.
  • tricalcium phosphate releases acidic degradation products.
  • some embodiments may include a composite including a bioceramic having acidic degradation products upon exposure to bodily fluids.
  • the acidic degradation products can increase the degradation rate of the polymer which can decrease the degradation time of the device.
  • a composite can have bioceramic particles that have basic degradation products.
  • hydroxyapatite releases basic degradation products.
  • the basic degradation products of the bioceramic particles can reduce the autocatalytic effect of the polymer degradation by neutralizing the acidic degradation products of the polymer degradation.
  • the basic degradation products of the bioceramic particles can reduce the degradation rate of the polymer.
  • bioceramic particles having a basic degradation product may also depress foreign body reaction.
  • bioceramic particles having a basic degradation product may buffer the reaction and neutralize the local pH drop.
  • some semi-crystalline biodegradable polymers have a degradation rate that is slower than desired for certain stent treatments.
  • the degradation time of a stent made from such polymers can be longer than desired.
  • a stent made from poly(L-lactide) can have a degradation time of between about two and three years. In some treatment situations, a degradation time of less than a year may be desirable, for example, between four and eight months.
  • the degradation of a hydrolytically degradable polymer follows a sequence including water penetration into the polymer followed by hydrolysis of bonds in the polymer.
  • the degradation of a polymer can be influenced by its affinity for water and the diffusion rate of water through the polymer.
  • a hydrophobic polymer has a low affinity for water which results in a relatively low water penetration.
  • the diffusion rate of water through crystalline regions of a polymer is lower than amorphous regions.
  • a biodegradable implantable medical device may be fabricated from a copolymer.
  • the copolymer can be a matrix in a bioceramic/polymer composite.
  • the copolymer can include hydrolytically degradable monomers or functional groups that provide desired degradation characteristics.
  • the copolymer can include functional groups that increase water penetration and water content of the copolymer.
  • a copolymer can include a primary functional group and at least one additional secondary functional group.
  • the copolymer may be a random copolymer including the primary functional group and at least one additional secondary functional group.
  • the copolymer with at least one secondary functional group can have a higher degradation rate than a homopolymer composed of the primary functional group.
  • a stent fabricated from the copolymer can have a lower degradation time than a stent fabricated from a homopolymer composed of the primary functional group.
  • the weight percent of the secondary functional group can be selected or adjusted to obtain a desired degradation rate of the copolymer or degradation time of a stent made from the copolymer.
  • the weight percent of the secondary functional group in a copolymer can be at least 1%, 5%, 10%, 15%, 30%, 40%, or, at least 50%.
  • a secondary functional group can be selected and the weight percent of the a secondary functional group can be adjusted so that the degradation time of a stent, with or without dispersed bioceramic particles, can be less than 18 months, 12 months, 8 months, 5 months, or more narrowly, less than 3 month.
  • the copolymer with at least one secondary functional group can have a lower crystallinity than a homopolymer composed of the primary functional group. It is believed that inclusion of a secondary functional group can perturb the crystalline structure of a polymer including the primary functional group, resulting in a reduced crystallinity. As a result, a stent fabricated from the copolymer has a larger percentage of amorphous regions, which allow greater water penetration. Thus, the degradation rate of the copolymer can be increased and the degradation time of a stent made from the copolymer can be decreased.
  • a secondary functional group can be a stereoisomer of the primary functional group.
  • One exemplary embodiment can include a poly(L-lactide-co-DL-lactide) copolymer.
  • DL-lactide can be a secondary functional group that perturbs the crystalline structure of the poly(L-lactide) so that the copolymer has a lower crystallinity than the poly(L-lactide).
  • increasing the number of secondary functional groups in the copolymer can result in a decrease in modulus of the copolymer as compared to a homopolymer of the primary functional group.
  • the decrease in modulus can be due to the decrease in crystallinity.
  • the decrease in modulus can reduce the ability of a stent to support a vessel.
  • the weight percent of the secondary functional group can be adjusted so that the ability of a stent to act as structural support is not substantially reduced.
  • the inclusion of bioceramic particles in the copolymer can partially or completely compensate for the reduction in the modulus of the copolymer.
  • the copolymer can include at least one secondary functional group with a greater affinity for water than the primary functional group.
  • the secondary functional group can be less hydrophobic or more hydrophilic than the primary functional group.
  • the decreased hydrophobicity or increased hydrophilicity can increase the concentration of water near bonds prone to hydrolysis, increasing the degradation rate and lowering the degradation time of a stent made from the copolymer.
  • the secondary functional groups can be selected so that segments of the copolymer with a secondary functional group can degrade faster than the primary functional group segments.
  • the difference in degradation rate can be due to the secondary functional groups being more hydrolytically active than the primary functional group.
  • a secondary functional group can be selected such that a homopolymer including the secondary functional group has a higher degradation rate than a homopolymer including the primary functional group.
  • the copolymer can be poly(L-lactide-co-glycolide).
  • the primary functional group can be L-lactide and the secondary functional group can be glycolide.
  • the weight percent of the glycolide in the copolymer can be at least 1%, 5%, 10%, 15%, 30%, 40%, or, at least 50%. In certain exemplary embodiments, the weight percent of glycolide group can be adjusted so that the degradation time of a stent, with or without dispersed bioceramic particles, can be less than 18 months, 12 months, 8 months, 5 months, or more narrowly, 3 months or less.
  • a bioceramic/polymer composite and fabrication of an implantable medical device therefrom.
  • a composite of a polymer and bioceramic particles can be extruded to form a polymer construct, such as a tube.
  • a stent can then be fabricated from the tube.
  • the composite can be formed in a number of ways.
  • the composite can be formed by melt blending. In melt blending the bioceramic particles are mixed with a polymer melt. The particles can be mixed with the polymer melt using extrusion or batch processing.
  • the bioceramic particles can be combined with a polymer in a powdered or granular form prior to melting of the polymer.
  • the particles and polymer can be mixed using mechanical mixing or stirring such as agitation of the particles and polymer in a container or a mixer.
  • the agitated mixture can then be heated to a temperature above the melt temperature of the polymer in an extruder or using batch processing.
  • a problem with the mechanical mixing or stirring techniques is that the polymer and particles may be separated into separate regions or layers. This is particularly a problem with respect to smaller particles such as nanoparticles.
  • obtaining a uniform dispersion by mixing particles with a polymer melt as described is that particles can agglomerate or form clusters.
  • the agglomeration of bioceramic particles makes it difficult to disperse the particles within the composite.
  • the presence of larger clusters in the composite tends to result in a decrease in material performance.
  • Such larger clusters can result in the formation of voids in a composite device, which are preferential sites for crack initiation and failure.
  • the mechanical mixing in a conventional single screw extruder or in batch processing can be insufficient to break up the clusters, resulting in a nonuniform mixture of bioceramic particles and polymer.
  • Various embodiments of forming a composite may be employed to increase the uniformity of dispersion of bioceramic particles within a polymer in a composite.
  • One set of embodiments may include forming a composite from a suspension of bioceramic particles and a polymer solution.
  • a composite formed using a suspension may result in a composite having more uniformly dispersed particles than the mixing methods described above.
  • Another set of embodiments can include reducing the agglomeration of bioceramic particles by decreasing the surface energy between particles to improve the dispersion of particles in the composite.
  • the surface energy can be decreased by treating the bioceramic particles with a surface modifier.
  • agglomeration can be reduced through mechanical mixing that applies shear stress to the particles sufficient to reduce the size of clusters of particles.
  • bioceramic particles can be mixed with a polymer by solution blending in which a composite mixture of bioceramic particles and polymer is formed from a suspension of particles in a polymer solution.
  • Certain embodiments of a method of forming an implantable medical device may include forming a suspension including a fluid, a polymer, and bioceramic particles.
  • a “suspension” is a mixture in which particles are suspended or dispersed in a fluid.
  • the fluid can be a solvent for the polymer so that the polymer is dissolved in the fluid.
  • the particles can be mixed with the fluid before or after dissolving the polymer in the fluid.
  • the suspension can be treated with ultrasound, for example, by an ultrasonic mixer.
  • the method may further include removing some or all of the fluid or separating the polymer and bioceramic particles from the fluid to obtain a composite including bioceramic particles dispersed within the polymer.
  • the polymer and bioceramic particles can be separated from the fluid by combining the suspension with a second fluid that may be a poor solvent for the polymer. At least some of the polymer may be allowed to precipitate upon combining the suspension solution with the second fluid. In some embodiments, at least some of the bioceramic particles may precipitate from the suspension with the precipitated polymer to form a composite mixture.
  • the precipitated composite mixture may then be filtered out of the solvents.
  • the filtered composite mixture can be dried to remove residual solvents.
  • the composite mixture can be dried in a vacuum oven or by blowing heated gas on the mixture.
  • Exemplary polymers may include, but are not limited to, poly (L-lactic acid), poly (DL-lactic acid), poly(lactide-coglycolide).
  • Representative solvents for such polymers can include toluene and chloroform.
  • Representative poor solvents for these polymers that may be used to precipitate the polymer include methanol, ethanol, isopropanol, and various alkanes such as hexane or heptane.
  • the particles can have strong interactions with polymer chains in solution which can result in particles becoming encapsulated or surrounded by polymer chains.
  • the interactions of the particles with the polymer can overcome interactions of the particles with the solution so that the particles precipitate with the polymer.
  • the degree of precipitation refers to the amount of particles that precipitate out of the suspension.
  • the degree of dispersion of particles within the precipitated polymer refers to the degree of mixing of the particles with the polymer.
  • the amount of polymer can be quantified by the weight percent of the polymer in the suspension solution.
  • the viscosity of the solution is also related to the amount of polymer in the solution. The higher the weight percent of dissolved polymer, the higher is the viscosity of the suspension solution.
  • the amount of particles precipitating can be relatively low.
  • the degree of dispersion of particles in the precipitated polymer tends to decrease. It is believed that at higher weight percent of polymer or higher viscosity, the interactions between polymer chains reduce the interaction of particles with polymer chains that cause particles to precipitate. For example, particles may be unable to move freely among the polymer chains.
  • a given suspension can have a particular combination of type of particles, particle concentration, and solvent.
  • the polymer weight percent or viscosity that can be varied to obtain both a desired degree of precipitation of particles and degree of dispersion of particles in the precipitated polymer.
  • the manner of combining the suspension with the poor solvent can also affect the degree of precipitation and degree of dispersion. For example, depositing a fine mist of small droplets into a poor solvent can more readily result in a desired degree of precipitation and degree of dispersion.
  • the manner of combining the suspension with the poor solvent can influence the range of polymer weight percent or viscosity that results in a desired degree of precipitation and degree of dispersion.
  • the polymer and bioceramic particles can be separated from the fluid in the suspension by evaporating the fluid in the suspension solution. In some embodiments, at least some of the bioceramic particles may be separated from the suspension solution along with the polymer upon evaporation to form the composite mixture. In one embodiment, the solvent can be removed by evaporation at high vacuum in a vacuum oven.
  • the composite formed from the suspension solution can be conveyed into an extruder.
  • the composite mixture may be extruded at a temperature above the melting temperature of the polymer and less than the melting temperature of the bioceramic particles.
  • the dried composite mixture may be broken into small pieces by, for example, chopping or grinding. Extruding smaller pieces of the composite mixture may lead to a more uniform distribution of the nanoparticles during the extrusion process.
  • the extruded composite mixture may then be formed into a polymer construct, such as a tube or sheet which can be rolled or bonded to form a tube.
  • a medical device may then be fabricated from the construct.
  • a stent can be fabricated from a tube by laser machining a pattern in to the tube.
  • a polymer construct may be formed from the composite mixture using an injection molding apparatus.
  • Preparation of a desired amount of precipitated composite mixture may require a large amount of solvent and precipitant. Therefore, in some embodiments, it may be advantageous to melt blend precipitated composite mixture with an amount of polymer in an extruder or in a batch process.
  • the polymer can be the same or a different polymer of the precipitated composite mixture. For example, a relatively small amount of precipitated composite mixture that has a weight percent of bioceramic particles higher than is desired can be prepared.
  • the precipitated composite mixture may be melt blended with an amount of biodegradable polymer to form a composite mixture than has a desired weight percent of bioceramic particles.
  • embodiments for increasing the dispersion of bioceramic particles in a composite can include decreasing the surface energy of particles with a surface modifier.
  • Bioceramic particles tend to aggregate due to high surface energy resulting from strong interparticle interactions.
  • Certain embodiments can include processing a plurality of bioceramic particles with a non-reactive surface modifier to reduce the surface energy between the particles, hence, agglomeration of particles.
  • a composite can then be formed with the processed bioceramic particles, the composite including the processed bioceramic particles dispersed within a polymer.
  • the processing can include forming a suspension of bioceramic particles including a suitable solvent and a non-reactive surface modifier dissolved in the solvent.
  • the non-reactive surface modifier tends to reduce surface energy between the particles, and thus, the fracture strength of the agglomerates or clusters of bioceramic particles.
  • the non-reactive surface modifier may cause or facilitate the break-up or reduction in size of the agglomerates or clusters of particles.
  • the agglomerates are more susceptible to break-up from application of conventional stirring techniques.
  • the breakdown of the agglomerates can be facilitated by mixing the suspension with an ultrasonic and/or high speed mixer known to one of skill in the art.
  • the suspension can be mixed to obtain a desired degree of break-up of agglomerates.
  • the suspension can be mixed from one to five minutes, five to 30 minutes, 30 minutes to an hour, one hour to eight hours, or more than eight hours.
  • the solvent may be removed and then mixed it with a polymer to form composite, e.g. by melt processing.
  • the polymer of the composite can be dissolved in the solvent after forming the suspension with the non-reactive surface modifier and nano bioceramic.
  • the polymer can be dissolved prior to adding the non-reactive surface modifier.
  • the solvent can be removed to form a composite mixture of polymer and bioceramic particles.
  • the solvent can be removed as described above through precipitation or through evaporation.
  • non-reactive surface modifiers for treating bioceramic particles include, but are not limited to, stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide (PEO-b-PPO-b-PEO), and polyethylene oxide-b-polylactide.
  • the non-reactive surface modifiers are used in amounts effective to prevent agglomeration of bioceramic particles.
  • the non-reactive surface modifiers can be used in amounts of about 0.1 weight % to about 3 weight % by weight of the total composite mixture, more narrowly about 0.3 weight % to about 2 weight %, or more narrowly about 0.5 weight % to about 1 weight %.
  • further embodiments for decreasing agglomeration and increasing the dispersion of bioceramic particles in a composite can include processing a mixture of particles and polymer with mechanical methods sufficient to reduce agglomeration.
  • Certain embodiments can include processing a mixture of a polymer and agglomerated bioceramic particles under high shear stress conditions.
  • Some embodiments can include processing the mixture such that the particles are subjected to shear stress higher than the fracture strength of the agglomerated particles.
  • a polymer melt can blended or mixed with bioceramic particles in a manner that subjects the mixture to a shear stress higher than the fracture strength of agglomerates of bioceramic particles.
  • polymer/bioceramic mixture can be processed so that a maximum shear stress generated during the melt blending is higher than the fracture strength of the bioceramic particle agglomerates. Agglomerated particles may be mechanically broken down and more uniformly dispersed within the polymer.
  • the mechanical mixing of polymer and bioceramic particles provided by a single-screw extruder is generally insufficient to obtain a uniformly dispersed mixture of bioceramic particles with the polymer.
  • the shear stress produced by a single screw extruder is typically lower than the fracture strength of bioceramic particle agglomerates.
  • mixing devices that can apply a shear stress higher than the fracture strength of agglomerates to a polymer melt.
  • Mechanical blending devices that can apply a sufficiently high shear stress include, but are not limited to, a twin screw extruder or batch mixer. During blending, once the shear stress is higher than the fracture strength of bioceramic particles agglomerates, the agglomerates are broken down and more uniformly dispersed into the polymer.
  • the polymer and bioceramic particles can be fed into a mechanical blending device separately and processed at high shear stress.
  • a composite mixture of polymer and bioceramic particles can be fed into a mechanical blending device and processed at the high shear stress.
  • the polymer/bioceramic particle mixture can be processed at the sufficiently high shear stress for a time sufficient to reduce agglomeration and disperse the particles.
  • the mixture can be processed between about 5 min. to about 30 min., more narrowly about 8 min. to about 20 min., or more narrowly about 10 min to about 15 min.
  • a composite of polymer and bioceramic particles formed by other methods disclosed herein can be processed at the high shear stress conditions.
  • the composite formed from a suspension solution through evaporation or precipitation can be processed at high shear stress conditions, for example, in a twin screw extruder or a batch mixer.
  • the composite formed with surface modified bioceramic particles can also be processed in this manner.
  • bioceramic particles treated with a non-reactive surface modifier can be processed with a polymer to form a composite.
  • the bioceramic particles can be treated by dispersing the bioceramic particles in a solution including the non-reactive surface modifier. Upon removal of the solvent from the solution, the treated particles can be processed in a mechanical mixing device such as an extruder. Since the fracture strength of clusters of agglomerated particles is decreased by the non-reactive surface modifier, even a conventional single screw extruder may facilitate break-up of the clusters, reducing agglomeration.
  • the treated particles and polymer can be processed in a device, such as a twin screw extruder or batch mixer, that subjects the particles and polymer to shear stress to shear stress greater than the fracture strength of clusters of untreated bioceramic particles.
  • a device such as a twin screw extruder or batch mixer
  • PEO/PLA polyphosphazenes
  • biomolecules such as fibrin, fibrinogen, cellulose, starch, collagen and hyaluronic acid
  • polyurethanes silicones
  • polyesters polyolefins, polyisobutylene and ethylene-alphaolefin copolymers
  • acrylic polymers and copolymers other than polyacrylates vinyl halide polymers and copolymers (such as polyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether), polyvinylidene halides (such as polyvinylidene chloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such as polystyrene), polyvinyl esters (such as polyvinyl acetate), acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon 66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,
  • polymers that may be especially well suited for use in fabricating an implantable medical device according to the methods disclosed herein include ethylene vinyl alcohol copolymer (commonly known by the generic name EVOH or by the trade name EVAL), poly(butyl methacrylate), poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508, available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride (otherwise known as KYNAR, available from ATOFINA Chemicals, Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethylene glycol.
  • EVAL ethylene vinyl alcohol copolymer
  • poly(vinylidene fluoride-co-hexafluororpropene) e.g., SOLEF 21508, available from Solvay Solexis PVDF, Thorofare, N.J.
  • hydroxyapatite nano particles are used as the bioceramic nanoparticles.
  • calcium sulfate nanoparticles are used as the nanoparticles.
  • the polymers used in the Examples were poly(L-lactic acid) (PLLA), poly(DL-lactic acid) (PDLLA), and poly(lactic acid-co-glycolide) (PLGA).
  • Step 1 Add bioceramic particles into suitable solvent, such as chloroform, acetone, etc. and stir to form a bioceramic particle suspension solution.
  • suitable solvent such as chloroform, acetone, etc.
  • Step 2 Slowly add a polymer such as PLLA, PDLLA, PLGA into suspension solution and stir until polymer dissolves completely.
  • the solution may still have a relatively low viscosity.
  • the bioceramic particles should be well dispersed while stirring.
  • Step 3 Slowly add the polymer into solution again to gradually increase solution viscosity. Repeat this step as needed until the polymer is completely dissolved and reasonable solution viscosity is developed.
  • Step 4 Apply ultrasonic mixing to suspension solution for 15-30 min to further disperse all the HAP uniformly into the PLLA solution.
  • Step 6 Add suspension solution to methanol to precipitate polymer and particles.
  • Step 1 Added 50 mg HAP particles into 300 mL of chloroform and stirred for 10-30 minutes to form bioceramic particle suspension solution.
  • Step 2 Slowly added 5 g PLLA into suspension solution and stirred about 8 hours to dissolve all polymer.
  • Step 3 Applied ultrasonic mixing to suspension solution for 15-30 min to further disperse HAP particles into PLLA solution.
  • Step 4 Added suspension solution to 1 L methanol to precipitate polymer and particles.
  • Step 5 Filtered the precipitate and dried about 24 hours in vacuum oven at 60° C. End product is PLLA/HAP composite.
  • Tensile testing of the composite samples and a pure PLLA were performed using an Instron tensile tester obtained from Instron in Canton, Mass. Test samples were prepared by hot pressing the PLLA/HAP composites and pure polymer to a thin film at 193° C. for 60 seconds. Testing bars were cut from the thin film and tested. The peak stress at break, the strain at break, and the Young's modulus were measured. The draw rate was about 0.5 in/min.
  • FIGS. 4-6 illustrate tensile testing results for pure PLLA and PLLA/HAP composites with 1 wt % HAP.
  • FIG. 4 depicts the peak stress for the two samples. PLLA/HAP composites have a higher peak stress than the pure PLLA.
  • FIG. 5 depicts the % strain at break for the two samples. The 1 wt % composite had a higher % strain at break than the pure PLLA.
  • FIG. 6 depicts Young's modulus for the two samples. The 1 wt % samples had a higher Young's modulus than the pure PLLA.
  • Step 1 Added 25 g HAP particles to 3 L chloroform and stirred for 10-30 minutes to form bioceramic particle suspension solution.
  • Step 2 Slowly added 50 g PLLA into suspension solution and stirred about 8 hours to dissolve all polymer.
  • Step 3 Applied ultrasonic mixing for 15-30 min to further disperse HAP particles into PLLA solution.
  • Step 4 Added suspension solution to 9 L methanol to precipitate particles and polymer.
  • Step 5 Filtered the precipitate and dried about 24 hours in vacuum oven at 60° C. End product is PLLA/HAP composite.
  • Step 1 Broke 2:1 wt/wt composite into small pieces
  • Step 2 Mixed 24 g of broken up composite and 376 g PLLA.
  • Step 3 Extruded mixture at 216° C.
  • Step 1 Broke PLGA/HAP composite into small pieces
  • Step 2 Mixed 12 g broken up PLGA/HAP composite and 396 g PLGA.
  • Step 3 Extruded mixture at 216° C.
  • the interfacial adhesion can be enhanced.
  • the adhesion between bioceramic particles and a biodegradable polymer can be improved by coating at least a portion of the surfaces of the bioceramic particles with an adhesion promoter such as 3-aminopropyltrimethoxysilane and 3-aminopropyltriethoxysilane.
  • Step 1 Added 100 ml distillated water to 1900 ml Ethanol and stirred for 15-30 min.
  • Step 2 Added 20 g 3-aminopropyltrimethoxysilane to water-ethanol mixture and stirred for 1 h.
  • Step 3 Added 20 g HAP and stirred for 2 h.
  • Step 4 Centrifuged the modified HAP from solution.
  • Step 5 Dried HAP about 8 hours in vacuum oven.
  • a stent was fabricated from PLGA/nanoparticle composite tubing. Prior to cutting a stent pattern, the tubing was expanded at 109° C. in a blow molder to increase radial strength. A stent pattern was cut in the expanded tubing using an ultra-fast pulse laser. The stent was crimped at 30° C. After crimping, the stent was cold sterilized.
  • FIGS. 7 and 8 illustrate compression testing results for PLGA/HAP stent (100:1 wt/wt). The compression testing results for 100% PLGA is also included for comparison.
  • FIG. 7 shows that the average radial strength of PLGA/HAP stent is about 11% higher than that of the 100% PLGA stent.
  • FIG. 8 shows that the compression modulus of PLGA/nano HAP stent increased by about 23% over the 100% PLGA stent.
  • FIG. 9 shows the recoil of PLGA/nano HAP stent is about 15% less than the PLGA stent.
  • Stents were fabricated from a bioceramic/polymer composite with a matrix of PLGA copolymer. Calcium sulfate nanoparticles, which were pretreated by PEG-PPG-PEG surface modifier, were mixed with the copolymer matrix.
  • PEG refers to polyethylene glycol
  • PPG refers to polypropylene glycol.
  • the copolymer had 85 wt % L-lactide and 15 wt % GA.
  • the composite was fabricated according to methods described herein.
  • the stents were fabricated from tubes made from the composite and the tubes was radially expanded. The expanded tubes were laser cut to form stents.
  • Stents were fabricated and tested having a ratio of copolymer to particles of 100:1 by weight. Five stent samples were used for each testing (compression, recoil or expansion testing). at a zero time point (after fabrication) and after 16 hours of accelerated aging. Accelerated aging refers to aging at 40° C.
  • the zero time point samples were subjected to compression testing. Compression tests were performed with an Instron testing machine obtained from Instron in Canton, Mass. A stent sample was placed between two flat plates. The plates were adjusted so that the distance between the plates was the diameter of the stent in an uncompressed state. The plates were then adjusted to compress the stent by 10%, 15%, 25%, and 50% of the uncompressed plate distance. A resistance force, the amount force in units of Newtons/mm that was necessary to keep the stent at each compression distance, was measured. The resistance force corresponds to a measure of the radial strength (Rs). The modulus was also measured. Table 1 provides the compression test results for the 100:1 samples at zero time point. The Rs and modulus correspond to 50% compression.
  • the stent samples were subjected to recoil testing at both time points for each bioceramic composition.
  • Each stent sample was deployed inside a length of Tecoflex elastic tubing.
  • Tecoflex tubing can be obtained from Noveon, Inc., Cleveland, Ohio.
  • the inside diameter (ID) of the Tecoflex tubing was 0.118 in and the outside diameter (OD) was 0.134 in.
  • ID of the sheath to hold the stent after crimping was 0.053 in.
  • the stents were deployed to an outer diameter of 3.0 mm by inflating the balloon. The balloon was deflated, allowing the stent to recoil. The diameter of the recoiled stent was then measured.
  • % Recoil (Inflated Diameter ⁇ Deflated Diameter)/Inflated Diameter ⁇ 100%
  • the Inflated Diameter is the diameter of a deployed stent prior to deflating the balloon and the Deflated Diameter is the diameter of the recoiled stent.
  • the deployment of the stent also results in an increase in length of the stent.
  • Table 2 provides the results of the recoil test for the five 100:1 stent samples at zero time point.
  • the outside diameter (OD) of the stent when the balloon was inflated and deflated was measured at a proximal end, a middle, and distal end of the stent.
  • ODwt refers to the outside diameter with tecoflex tubing
  • ODS refers to the outside diameter with stent only.
  • the % Total Recoil given is the average of the recoil calculated at each point along the length of the stent. TABLE 2 Recoil test results for 100:1 stent samples at zero time point.
  • the stents were subject to expansion testing at both time points. Each stent sample was deployed to an outer diameter of first 3.0 mm, then 3.5 mm, and then 4.0 mm. The number of cracks in the stents were counted at each deployment diameter.
  • Table 3 provides the number of cracks observed in the 100:1 stent samples at zero time point when deployed at a diameter of 3.5 mm.
  • the columns refer to the size of the cracks: “Micro” refers to micro-sized cracks, “ ⁇ 25%” refers to cracks less than 25% of the strut width, “ ⁇ 50%” refers to cracks less than 50% of the strut width, and “>50%” refers to cracks greater than 50% of the strut width.
  • “A” refers to the “v-shaped” region or bending region of struts and “B” refers to a spider region at the intersection of 5 struts. The regions can be seen in FIGS. 10 A-D.
  • FIGS. 10 A-D depict photographic images of the 100:1 stent sample 1 at zero time point before expansion, expanded to 3.5 mm, after recoil, and after expansion to 4.0 mm, respectively. TABLE 3 Number of cracks observed at 3.5 mm deployment for 100:1 stent samples at zero time point. Stents Deployed to 3.5 mm Stent # Micro ⁇ 25% ⁇ 50% >50% 1 A 4 2 B 2 A 2 B 3 A 4 B 4 A B 5 A 1 B
  • Table 4 provides the results of the recoil test for the 100:1 stent samples at 16 hours of accelerated aging.
  • Tables 5A-B provide the number of cracks observed in the 100:1 stent samples at 16 hours of accelerated aging when deployed at a diameter of 3.5 mm and 4.0 mm, respectively. No broken struts were observed at either deployment diameter.
  • FIGS. 11 A-D depict photographic images of the 100:1 stent sample 1 at 16 hours of accelerated aging before expansion, expanded to 3.5 mm, after recoil, and after expansion to 4.0 mm, respectively. TABLE 4 Recoil test results for 100:1 stent samples at 16 hours of accelerated aging.
  • a stent was fabricated from a polymer/bioceramic composite.
  • the matrix was a PLGA copolymer. Calcium sulfate nanoparticles were mixed with the copolymer matrix.
  • the copolymer was 50 wt % L-lactide and 50 wt % glycolide. The ratio of copolymer to particles was 100:3 by weight.
  • the composite was fabricated according to methods described herein. A tube was fabricated from the composite and the tube was radially expanded. The expanded tube was laser cut to form a stent. The expected degradation time for the composite stent is about 3-5 months.
  • the following example illustrates formation of a polymer/bioceramic particle composite with surface modified bioceramic particles.
  • Step 1 A solution was formed by dissolving 8 g stearic acid (surface modifier) in 200 ml isopropanol (solvent).
  • Step 2 20 g Calcium sulfate (CS) nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • CS Calcium sulfate
  • Step 3 The suspension was subjected to ultrasonic mixing for 30-60 min to reduce the agglomeration of nanoparticles.
  • Step 4 The solvent was removed through evaporation at high vacuum.
  • Step 5 A polymer/nanoparticle composite was prepared by blending two portions of the modified calcium sulfate nanoparticles and 100 portions of a PLLA composite mixture which contained uniformly dispersed nanoparticles in the PLLA matrix.
  • Step 1 A solution was formed by dissolving 20 g PEO-b-PPO-b-PPO (surface modifier) in 400 ml chloroform (solvent).
  • Step 2 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3 The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4 A polymer/nanoparticle composite was prepared by adding 20 g LPLG to the modified calcium sulfate nanoparticles. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 5 The solvent was removed through evaporation at high vacuum.
  • Step 1 A solution was formed by dissolving 20 g PEO-b-PPO-b-PPO (surface modifier) in 400 ml chloroform (solvent).
  • Step 2 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3 The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4 The solvent was removed through evaporation at high vacuum.
  • Step 5 A polymer/nanoparticle composite was prepared by mixing the modified calcium sulfate nanoparticles with 1000 g LPLG. The tubing was prepared through extrusion using single screw extruder.
  • FIG. 12A shows s SEM photographic image of an LPLG/unmodified CS particle composite.
  • the CS particles size range from 1 to 50 ⁇ m, with most particles between 2 to 30 ⁇ m.
  • FIG. 12B illustrates a photographic image of an LPLG/modified CS particle composite with particles modified with PEO-b-PPO-b-PEO.
  • the CS particles size range from 0.1 to 10 ⁇ m, where most of the particles are from 0.3 to 6 ⁇ m.
  • Step 1 A solution was formed by dissolving 16 g stearic acid (surface modifier) in 400 ml chloroform.
  • Step 2 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3 The suspension was subjected to ultrasonic mixing for 30 min.
  • Step 5 A polymer/nanoparticle composite was prepared by adding 20 g LPLG to the modified calcium sulfate nanoparticles. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 4 The solvent was removed through evaporation at high vacuum.
  • Step 1 A solution was formed by dissolving 16 g stearic acid (surface modifier) in 400 ml chloroform (solvent).
  • Step 2 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3 The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4 The solvent was removed through evaporation at high vacuum.
  • Step 5 A polymer/nanoparticle composite was prepared by mixing the modified calcium sulfate nanoparticles with 1000 g LPLG. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 6 The cross-section morphology of the tubing was observed by SEM ( FIG. 1C ).
  • FIG. 12C shows a photographic image of LPLG/stearic acid modified CS particle composite, which is similar to the morphology of LPLG/PEO-b-PPO-b-PEO modified CS particle composite.
  • the CS particles sizes range from 0.1 to 10 ⁇ m, where most of them are from 0.5 to 6 ⁇ m.
  • LPLG material and surface modified or non-modified CS nano particles were mixed and extruded through a twin screw extruder.
  • CS content in the extruded material is roughly 2%.
  • the morphology of the extruded material was observed by SEM
  • FIG. 13A illustrates a photographic image of extruded LPLG/CS composite using a twin screw extruder. Overall, the CS particles sizes range from 0.2 to 20 ⁇ m with most of the particles from 0.5 to 5 ⁇ m. As a comparison, FIG. 13B shows a photographic image of LPLG/CS particle composite using a single screw extruder with low shear stress. Overall, the CS particles sizes range from 1 to 50 ⁇ m, where most of them are from 2 to 30 ⁇ m. FIG.
  • FIG. 13C shows an image of an LPLG/PEO-b-PPO-b-PEO modified CS particles composite by twin screw extruder at with high shear stress, the CS particles sizes range from 0.1 to 10 ⁇ m, where most of them are from 0.2 to 3 ⁇ m.

Abstract

Methods and devices relating to polymer-bioceramic composite implantable medical devices are disclosed.

Description

    CROSS-REFERENCE
  • This application is a continuation-in-part of U.S. application Ser. No. 11/706,550 filed on Feb. 14, 2007, and a continuation-in-part of U.S. application Ser. No. 11/529,996 filed Sep. 29, 2006, and a continuation-in-part of U.S. application Ser. No. 11/523,866 filed on Sep. 19, 2006, which is a continuation-in-part of U.S. application Ser. No. 11/443,870 filed on May 30, 2006, which are all incorporated herein by reference.
  • BACKGROUND OF THE INVENTION
  • 1. Field of the Invention
  • This invention relates to implantable medical devices and methods of fabricating implantable medical devices.
  • 2. Description of the State of the Art
  • This invention relates to radially expandable endoprostheses, which are adapted to be implanted in a bodily lumen. An “endoprosthesis” corresponds to an artificial device that is placed inside the body. A “lumen” refers to a cavity of a tubular organ such as a blood vessel.
  • A stent is an example of such an endoprosthesis. Stents are generally cylindrically shaped devices, which function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce body vessels and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.
  • The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through a bodily lumen to a region, such as a lesion, in a vessel that requires treatment. “Deployment” corresponds to the expanding of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into a bodily lumen, advancing the catheter in the bodily lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.
  • In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves compressing or crimping the stent onto the balloon. The stent is then expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn. In the case of a self-expanding stent, the stent may be secured to the catheter via a constraining member such as a retractable sheath or a sock. When the stent is in a desired bodily location, the sheath may be withdrawn which allows the stent to self-expand.
  • The stent must be able to satisfy a number of mechanical requirements. First, the stent must be capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of a vessel. Therefore, a stent must possess adequate radial strength. Radial strength, which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around a circumferential direction of the stent. Radial strength and rigidity, therefore, may also be described as, hoop or circumferential strength and rigidity.
  • Once expanded, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. For example, a radially directed force may tend to cause a stent to recoil inward. Generally, it is desirable to minimize recoil. In addition, the stent must possess sufficient flexibility to allow for crimping, expansion, and cyclic loading. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure. Finally, the stent must be biocompatible so as not to trigger any adverse vascular responses.
  • The structure of a stent is typically composed of scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts or bar arms. The scaffolding can be formed from wires, tubes, or sheets of material rolled into a cylindrical shape. The scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment). A conventional stent is allowed to expand and contract through movement of individual structural elements of a pattern with respect to each other.
  • Additionally, a medicated stent may be fabricated by coating the surface of either a metallic or polymeric scaffolding with a polymeric carrier that includes an active or bioactive agent or drug. Polymeric scaffolding may also serve as a carrier of an active agent or drug.
  • Furthermore, it may be desirable for a stent to be biodegradable. In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Therefore, stents fabricated from biodegradable, bioabsorbable, and/or bioerodable materials such as bioabsorbable polymers should be configured to completely erode only after the clinical need for them has ended.
  • A potential problem with polymeric stents is that their struts or bar arms can crack during crimping and expansion. This is especially the case with brittle polymers. The localized portions of the stent pattern subjected to substantial deformation during crimping and expansion tend to be the most vulnerable to failure.
  • Therefore, it is desirable for a stent to have flexibility and resistance to cracking during deployment. It is also advantageous for a stent to be rigid and resistant to creep after deployment. It would also be desirable to be able to control the degradation rate of the device.
  • SUMMARY OF THE INVENTION
  • Certain embodiments of the present invention include a method of fabricating an implantable medical device comprising: treating the surface of a plurality of bioceramic particles with a non-reactive surface modifier; forming a composite including the treated bioceramic particles dispersed within a polymer, the non-reactive surface modifier reducing the agglomeration of the treated bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • Further embodiments of the present invention include a method for fabricating an implantable medical device comprising: forming a suspension solution including a fluid, a non-reactive surface modifier, a polymer and bioceramic particles, wherein the polymer is dissolved in the fluid, and wherein the bioceramic particles are dispersed in the solution; removing all or substantially all of the fluid to form a composite mixture, wherein the composite mixture comprises the bioceramic particles dispersed within the polymer, wherein the non-reactive surface modifier reduces the agglomeration of the bioceramic particles in the suspension solution and during formation of the composite mixture; and fabricating an implantable medical device from the composite mixture.
  • Additional embodiments of the present invention include a stent comprising: a structural element including a bioceramic/polymer composite, the composite having plurality of bioceramic particles dispersed within a polymer, a non-reactive surface modifier being on a surface of at least some of the particles, wherein the surface modifier reduces agglomeration of the bioceramic particles during the formation of the composite.
  • Certain other embodiments of the present invention include a method of fabricating an implantable medical device comprising: subjecting a mixture of bioceramic particles and polymer to high sheer stress using twin-screw extruder or a batch mixer; forming a composite including the bioceramic particles dispersed within a polymer, the mechanical smash, reducing the agglomeration of the bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • Some embodiments of the present invention include a method of fabricating an implantable medical device comprising: treating the surface of a plurality of bioceramic particles with a non-reactive surface modifier; subjecting a mixture of the treated bioceramic particles and polymer to high sheer stress using twin-screw extruder or a batch mixer; forming a composite including the treated bioceramic particles dispersed within a polymer, the non-reactive surface modifier or mechanical smash reducing the agglomeration of the treated bioceramic particles during the formation; and fabricating an implantable medical device from the composite.
  • BRIEF DESCRIPTION OF THE DRAWINGS
  • FIG. 1 depicts a three-dimensional view of a stent.
  • FIG. 2A depicts a section of a structural element from the stent depicted in FIG. 1.
  • FIG. 2B depicts bioceramic particles dispersed within a polymer matrix.
  • FIG. 3 depicts a schematic plot of the crystal nucleation rate, the crystal growth rate, and the overall rate of crystallization for a semicrystalline polymer.
  • FIG. 4 is a graph depicting the peak stress for PLLA and PLLA/HAP composites.
  • FIG. 5 is a graph depicting the break strain for PLLA and PLLA/HAP composites.
  • FIG. 6 is a graph depicting Young's modulus for PLLA and PLLA/HAP composite stent in compression testing.
  • FIG. 7 is a graph depicting the radial stress for PLGA and PLGA/HAP composite stent in compression testing.
  • FIG. 8 is a graph depicting the compression modulus for PLGA and PLGA/HAP composite stent in compression testing.
  • FIG. 9 is a graph depicting recoil testing of PLGA and PLGA/HAP composite stent.
  • FIGS. 10A-D depict photographic images of a PLGA/calcium sulfate composite stent with a weight ratio of PLGA/calcium sulfate of 100:1 at zero time point.
  • FIGS. 11A-D depict photographic images of a PLGA/calcium sulfate composite stent with a weight ratio of PLGA/calcium sulfate of 100:1 at 16 hours of accelerated aging.
  • FIG. 12A depicts a SEM photographic image of LPLG/unmodified calcium sulfate composite before surface modification
  • FIG. 12B depicts a SEM photographic image of LPLG/PEO-b-PPO-b-PEO modified calcium sulfate composite after surface modification
  • FIG. 12C depicts a SEM photographic image of LPLG/stearic acid modified calcium sulfate composite after surface modification
  • FIG. 13A depicts a photographic image of LPLG/calcium sulfate composite formed using twin screw extruder
  • FIG. 13B depicts a photographic image of LPLG/calcium sulfate composite formed using single screw extruder
  • FIG. 13C depicts a photographic image of LPLG/PEO-b-PPO-b-PEO modified calcium sulfate composite formed using twin screw extruder
  • DETAILED DESCRIPTION OF THE INVENTION
  • Those of ordinary skill in the art will realize that the following description of the invention is illustrative only and not in any way limiting. Other embodiments of the invention will readily suggest themselves to such skilled persons based on the disclosure herein. All such embodiments are within the scope of this invention.
  • For the purposes of the present invention, the following terms and definitions apply:
  • The “glass transition temperature,” Tg, is the temperature at which the amorphous domains of a polymer change from a brittle vitreous state to a solid deformable or ductile state at atmospheric pressure. In other words, the Tg corresponds to the temperature where the onset of segmental motion in the chains of the polymer occurs. When an amorphous or semicrystalline polymer is exposed to an increasing temperature, the coefficient of expansion and the heat capacity of the polymer both increase as the temperature is raised, indicating increased molecular motion. As the temperature is raised the actual molecular volume in the sample remains constant, and so a higher coefficient of expansion points to an increase in free volume associated with the system and therefore increased freedom for the molecules to move. The increasing heat capacity corresponds to an increase in heat dissipation through movement. Tg of a given polymer can be dependent on the heating rate and can be influenced by the thermal history of the polymer. Furthermore, the chemical structure of the polymer heavily influences the glass transition by affecting mobility.
  • “Stress” refers to force per unit area, as in the force acting through a small area within a plane. Stress can be divided into components, normal and parallel to the plane, called normal stress and shear stress, respectively. True stress denotes the stress where force and area are measured at the same time. Conventional stress, as applied to tension and compression tests, is force divided by the original gauge length.
  • “Strength” refers to the maximum stress along an axis which a material will withstand prior to fracture. The ultimate strength is calculated from the maximum load applied during the test divided by the original cross-sectional area.
  • “Modulus” may be defined as the ratio of a component of stress or force per unit area applied to a material divided by the strain along an axis of applied force that results from the applied force. For example, a material has both a tensile and a compressive modulus. A material with a relatively high modulus tends to be stiff or rigid. Conversely, a material with a relatively low modulus tends to be flexible. The modulus of a material depends on the molecular composition and structure, temperature of the material, amount of deformation, and the strain rate or rate of deformation. For example, below its Tg, a polymer tends to be brittle with a high modulus. As the temperature of a polymer is increased from below to above its Tg, its modulus decreases.
  • “Strain” refers to the amount of elongation or compression that occurs in a material at a given stress or load.
  • “Elongation” may be defined as the increase in length in a material which occurs when subjected to stress. It is typically expressed as a percentage of the original length.
  • “Toughness” is the amount of energy absorbed prior to fracture, or equivalently, the amount of work required to fracture a material. One measure of toughness is the area under a stress-strain curve from zero strain to the strain at fracture. Thus, a brittle material tends to have a relatively low toughness.
  • “Solvent” is defined as a substance capable of dissolving or dispersing one or more other substances or capable of at least partially dissolving or dispersing the substance(s) to form a uniformly dispersed solution at the molecular- or ionic-size level. The solvent should be capable of dissolving at least 0.1 mg of the polymer in 1 ml of the solvent, and more narrowly 0.5 mg in 1 ml at ambient temperature and ambient pressure.
  • As used herein, an “implantable medical device” includes, but is not limited to, self-expandable stents, balloon-expandable stents, stent-grafts, implantable cardiac pacemakers and defibrillators; leads and electrodes for the preceding; implantable organ stimulators such as nerve, bladder, sphincter and diaphragm stimulators, cochlear implants; prostheses, vascular grafts, grafts, artificial heart valves and cerebrospinal fluid shunts.
  • An implantable medical device can be designed for the localized delivery of a therapeutic agent. A medicated implantable medical device may be constructed by coating the device with a coating material containing a therapeutic agent. The substrate of the device may also contain a therapeutic agent.
  • FIG. 1 depicts a three-dimensional view of stent 100. In some embodiments, a stent may include a pattern or network of interconnecting structural elements 110. Stent 100 may be formed from a tube (not shown). Stent 100 includes a pattern of structural elements 110, which can take on a variety of patterns. The structural pattern of the device can be of virtually any design. The embodiments disclosed herein are not limited to stents or to the stent pattern illustrated in FIG. 1. The embodiments are easily applicable to other patterns and other devices. The variations in the structure of patterns are virtually unlimited. A stent such as stent 100 may be fabricated from a tube by forming a pattern with a technique such as laser cutting or chemical etching.
  • The geometry or shape of an implantable medical device may vary throughout its structure to allow radial expansion and compression. A pattern may include portions of structural elements or struts that are straight or relatively straight, an example being a portion 120. In addition, patterns may include structural elements or struts that include curved or bent portions such as portions 130, 140, and 150.
  • An implantable medical device can also be made partially or completely from a biodegradable, bioabsorbable, or biostable polymer. A polymer for use in fabricating an implantable medical device can be biostable, bioabsorbable, biodegradable or bioerodable. Biostable refers to polymers that are not biodegradable. The terms biodegradable, bioabsorbable, and bioerodable are used interchangeably and refer to polymers that are capable of being completely degraded and/or eroded when exposed to bodily fluids such as blood and can be gradually resorbed, absorbed, and/or eliminated by the body. The processes of breaking down and absorption of the polymer can be caused by, for example, hydrolysis and metabolic processes.
  • However, polymers tend to have a number of shortcomings for use as materials for implantable medical devices such as stents. Many biodegradable polymers have a relatively low modulus at the physiological conditions in the human body. In general, compared to metals, the strength to weight ratio of polymers is smaller than that of metals. A polymeric stent with inadequate radial strength can result in mechanical failure or recoil inward after implantation into a vessel. To compensate for the relatively low modulus, a polymeric stent requires significantly thicker struts than a metallic stent, which results in an undesirably large profile.
  • Another shortcoming of polymers is that many polymers, such as biodegradable polymers, tend to be brittle under physiological conditions or conditions within a human body. Specifically, such polymers can have a Tg above human body temperature which is approximately 37° C. These polymer systems exhibit a brittle fracture mechanism in which there is little or no plastic deformation prior to failure. As a result, a stent fabricated from such polymers can have insufficient toughness for the range of use of a stent.
  • Other potential problems with polymeric stents include creep, stress relaxation, and physical aging. Creep refers to the gradual deformation that occurs in a polymeric contruct subjected to an applied load. Creep occurs even when the applied load is constant.
  • It is believed that the delayed response of polymer chains to stress during deformation causes creep behavior. As a polymer is deformed, polymeric chains in an initial state rearrange to adopt a new equilibrium configuration. Rearrangement of chains takes place slowly with the chains retracting by folding back to their initial state. For example, an expanded stent can retract radially inward, reducing the effectiveness of a stent in maintaining desired vascular patency. The rate at which polymers creep depends not only on the load, but also on temperature. In general, a loaded construct creeps faster at higher temperatures.
  • Stress relaxation is also a consequence of delayed molecular motions as in creep. Contrary to creep, however, which is experienced when the load is constant, stress relaxation occurs when deformation (or strain) is constant and is manifested by a reduction in the force (stress) required to maintain a constant deformation.
  • Physical aging, as used herein, refers to densification in the amorphous regions of a semi-crystalline polymer. Densification is the increase in density of a material or region of a material. Densification, and thus physical aging, is also the result of relaxation or rearrangement of polymer chains.
  • Various embodiments of the present invention include an implantable medical device fabricated from a composite including a polymer matrix or continuous phase and bioceramic particles as a discrete phase. The bioceramic particles may tend to reduce or eliminate a number of the above-mentioned shortcomings of polymers. For example, the bioceramic particles can increase the toughness and modulus and modify the degradation rate of the polymer. In some embodiments, the composite may include a plurality of bioceramic particles dispersed within the polymer.
  • In general, it is desirable for the bioceramic particles to be uniformly dispersed throughout the biodegradable polymer. The more uniform the dispersion of the particles results in more uniform properties of the composite and a device fabricated from the composite. For example, a uniform dispersion can result in a uniform increase in toughness and modulus and modification of degradation rate. In some embodiments, the bioceramic particles are uniformly or substantially uniformly dispersed within the biodegradable polymer.
  • In certain embodiments, a structural element of an implantable medical device may be fabricated from a bioceramic/polymer composite. Structural elements can include, but are not limited to, any supporting element such as a strut, wire, or filament. FIG. 2A depicts a section 200 of a structural element 110 from stent 100. A portion 210 of section 200 is shown in an expanded view in FIG. 2B. FIG. 2B depicts bioceramic particles 220 dispersed throughout a polymer matrix 230.
  • Bioceramics can include any ceramic material that is compatible with the human body. More generally, bioceramic materials can include any type of compatible inorganic material or inorganic/organic hybrid material. Bioceramic materials can include, but are not limited to, alumina, zirconia, apatites, calcium phosphates, silica based glasses, or glass ceramics, and pyrolytic carbons. Bioceramic materials can be bioabsorbable and/or active. A bioceramic is active if it actively takes part in physiological processes. A bioceramic material can also be “inert,” meaning that the material does not absorb or degrade under physiological conditions of the human body and does not actively take part in physiological processes.
  • Illustrative examples of apatites and other calcium phosphates, include, but are not limited hydroxyapatite (Ca10(PO4)6(OH)2), floroapatite (Ca10(PO4)6F2), carbonate apatide (Ca10(PO4)6CO3), tricalcium phosphate (Ca3(PO4)2), octacalcium phosphate (Ca8H2(PO4)6-5H2O), octacalcium phosphate (Ca8H2(PO4)6-5H2O), calcium pyrophosphate (Ca2P2O7-2H2O), tetracalcium phosphate (Ca4P2O9), and dicalcium phosphate dehydrate (CaHPO4-2H2O).
  • The term bioceramics can also include bioactive glasses that are bioactive glass ceramics composed of compounds such as SiO2, Na2O, CaO, and P2O5. For example, a commercially available bioactive glass, Bioglass®, is derived from certain compositions of SiO2—Na2O—K2O—CaO—MgO—P2O5 systems. Some commercially available bioactive glasses include, but are not limited to:
  • 45S5: 46.1 mol % SiO2, 26.9 mol % CaO, 24.4 mol % Na2O and 2.5 mol % P2O5;
  • 58S: 60 mol % SiO2, 36 mol % CaO, and 4 mol % P2O5; and
  • S70C30: 70 mol % SiO2, 30 mol % CaO.
  • Another commercially available glass ceramic is A/W.
  • In some embodiments, bioceramic particles in a composite implantable medical device, may be used to inhibit or prevent infection since some bioceramics can have an anti-infective property. Bioceramics may release various ions such as calcium and phosphate ions which broadly exist in human body fluid and blood plasma. Examples of bioceramics that release calcium and/or phosphate ions include various calcium phosphates and bioactive glasses. The released ions may depress foreign body reaction. Trends Biomater. Artif. Tren, Vol 18 (1), pp 9-17.
  • As indicated above, an implantable medical device such as a stent can be medicated by incorporating an active agent in a coating over the device or within the substrate of the device. In some embodiments, the ions released from bioceramics can have an additive therapeutic and/or a synergistic therapeutic effect to the active agent. For example, ions can be used in conjunction with anti-proliferative and/or anti-inflammatory agents.
  • Bioceramic particles can be partially or completely made from a biodegradable, bioabsorbable, or biostable ceramic. Examples of bioabsorbable bioceramics include various types of bioglass materials, tetracalcium phosphate, amorphous calcium phosphate, alpha-tricalcium phosphate, and beta-tricalcium phosphate. Biostable bioceramics include alumina and zirconia.
  • Various sizes of the bioceramic particles may be used in the composite. For example, the bioceramic particles can include, but are not limited to, nanoparticles and/or micro particles. A nanoparticle refers to a particle with a characteristic length (e.g., diameter) in the range of about 1 nm to about 1,000 nm. A micro particle refers to a particle with a characteristic length in the range of greater than 1,000 nm and less than about 10 micrometers. Additionally, bioceramic particles can be of various shapes, including but not limited to, spheres and fibers.
  • Additionally, the particles size distribution can be important in modifying the properties of the polymer. Generally, a narrow size distribution is preferable.
  • The composite of a structural element of a device may have between 0.01% and 10% of bioceramic particles by weight, or more narrowly, between 0.5% and 2% bioceramic particles by weight as compared to the polymer matrix of the composite.
  • As indicated above, the bioceramic particles can reduce or eliminate a number of shortcomings of polymers that are used for implantable medical devices. In one aspect of the invention, bioceramic particles can increase the fracture toughness of polymers of implantable medical device. In general, the higher the fracture toughness, the more resistant a material is to the propagation of cracks. In some embodiments, bioceramic particles may be used in a composite having a matrix polymer that is brittle at physiological conditions. In particular, such a polymer can have a Tg above body temperature. In one embodiment, the bioceramic particles may be nanoparticles.
  • Certain regions of an implantable medical device, such as a stent, experience a high degree of stress and strain when the device is under stress during use. For example, when a stent is crimped and deployed, curved or bending regions such as portions 130, 140, and 150 can have highly concentrated strain which can lead to fracture. The bioceramic particles can increase fracture toughness by reducing the concentration of strain by dispersing the strain over a large volume of the material. Particles can absorb energy due to applied stress and disperse energy about a larger volume in the bioceramic/polymer composite.
  • Therefore, rather than being highly concentrated the stress and strain in a device fabricated from a bioceramic composite is divided into many small interactions involving numerous individual particles. When a crack is initiated in the material and starts traveling through the composite, the crack breaks up into finer and finer cracks due to interaction with the particles. Thus, the particles tend to dissipate the energy of imparted to the device by the applied stress. For a give weight ratio of particles to matrix, as the size of the particles decreases the number of particles dispersed throughout the device per unit volume also increases. Thus, the number of particles to disperse the energy of applied stress to the device increases. Therefore, it is advantageous to use nanoparticles to increase the toughness of the polymer. It has been shown that the fracture toughness of a polymeric material can be improved by using nanoparticles as a discrete or reinforcing phase in a composite. J. of Applied Polymer Science, 94 (2004) 796-802.
  • Bioceramic particles, more particularly nano-bioceramic particles, by providing more crystallites in a network in the bioceramic/polymer composite increase fracture toughness. In yet another aspect of the invention, bioceramic particles can be used to increase the the strength and modulus of the polymer. As indicated above, a polymeric stent requires a high radial strength in order to provide effective scaffolding of a vessel. Many biodegradable polymers have a relatively low strength and modulus as compared to metals. A composite with bioceramic particles with a higher modulus than a matrix polymer may have a higher strength and modulus than the polymer. The higher strength and modulus may allow for the manufacture of a composite stent with much thinner struts than a stent fabricated from the matrix polymer alone. Examples of relatively low modulus polymers include, but are not limited to, poly(D,L-lactide-co-glycolide), poly(lactide-co-caprolactone), poly(lactide-co-trimethylene carbonate), poly(glycolide-co-caprolactone), and poly(D,L-lactide). It has been reported that composites with nanoparticles can increase the modulus of a polymer by 1-2 orders of magnitude. Mechanical Properties of Polymers and Composites, Lawrence E. Nielsen and Robert F. Landel, 2nd ed., p. 384-385 (1993).
  • In addition, bioceramic particles in a polymer composite can also reduce or eliminate creep, stress relaxation, and physical aging. It is believed that particles can act as “net point” that reduce or inhibit movement of polymer chains in amorphous regions of a polymer.
  • Additionally, in composites fabricated from semicrystalline polymers, the crystallinity of a bioceramic/polymer composite that forms an implantable device can be controlled to reduce or eliminate creep, stress relaxation, and physical aging. As indicated above, these phenomena in a polymer are due to rearrangement or relaxation of polymer chains.
  • In general, as the crystallinity of a semicrystalline polymer increases, physical aging creep, and stress relaxation are reduced. This is likely due to the fact that polymer chains in the amorphous domains capable of movement are reduced by the crystalline domains. However, increasing crystallinity can result in brittleness in a polymer at physiological conditions.
  • In further embodiments, a structural element of an implantable medical device may include a composite having a plurality of crystalline domains dispersed within an amorphous biodegradable polymeric matrix phase. The crystalline domains may be formed around bioceramic particles. In certain embodiments, the composite that makes up the structural element may have a relatively low crystallinity. For example, the crystallinity can be less than 50%, 30%, 20%, or less than 10%.
  • Additionally, the device can be fabricated so that the resulting composite has a relatively large number of crystalline domains that are relatively small. In certain embodiments, the average crystal size can be less than 10 microns, or less than 5 microns. As the size of the crystalline domains decreases along with an increase in the number of domains, the polymer may become less brittle and, which increases the fracture toughness. Although the crystallinity of the resulting polymer can be relatively low, the presence of the relatively large number of relatively small crystalline domains can reduce or eliminate physical aging, creep, and stress relaxation.
  • The size and number of crystallites domains can be controlled during formation of a polymer construct from an implantable medical device is fabricated. Polymer constructs, such as tubes, can be formed using various types of forming methods, including, but not limited to extrusion or injection molding. Representative examples of extruders include, but are not limited to, single screw extruders, intermeshing co-rotating and counter-rotating twin-screw extruders, and other multiple screw masticating extruders.
  • In some embodiments, a mixture of a polymer and bioceramic particles can be extruded to form a polymer construct, such as a tube. A polymer melt mixed with the bioceramic particles can be conveyed through an extruder and forced through a die in the shape of as an annular film in the shape of a tube. The annular film can be cooled below the melting point, Tm, of the polymer to form an extruded polymeric tube. For example, the annular film may be conveyed through a water bath at a selected temperature. Alternatively, the annular film may be cooled by a gas at a selected temperature. The annular film may be cooled at or near an ambient temperature, e.g. 25° C. Alternatively, the annular film may be cooled at a temperature below ambient temperature.
  • In general, crystallization in a polymer tends to occur in a polymer at temperatures between Tg and Tm of the polymer. Therefore, in some embodiments, the temperature of the polymer construct during cooling can be between Tg and Tm. As the temperature of the extruded mixture is cooled below Tm to form a polymer construct, such as a tube, the bioceramic particles provide a point of nucleation in the polymer melt for the formation of crystalline domains.
  • A network of many small crystalline domains is formed, which can work to tie crystalline domains together and reduce, inhibit or prevent fracturing, creep, stress relaxation, and physical aging of the polymer. The crystalline domains can serve as net points in the amorphous domains that restrict the freedom of movement of polymer chains in the amorphous domain. As a result, physical aging, creep, and stress relaxation can be reduced. In addition, for the reasons discussed above, the toughness of the polymer is also increased.
  • In general, both microparticles and nanoparticles can be used as nucleation points. However, as the number of particles increases and size of the particles decreases, the crystalline domains become more effective in increasing fracture toughness and reducing physical aging, creep, and stress relaxation. The closer the crystalline domains are to one another within the amorphous domain of a polymer, the more the crystalline domains can limit the degree of freedom movement of polymer chains in the amorphous domain. Therefore, nanoparticles may be more effective in reducing physical aging, creep, and stress relaxation.
  • In certain embodiments, the size of the crystalline domains can be controlled by the temperature of the cooling polymer construct from an extruder. In general, crystallization tends to occur in a polymer at temperatures between Tg and Tm of the polymer. The rate of crystallization in this range varies with temperature. FIG. 3 depicts a schematic plot of the crystal nucleation rate (RN), the crystal growth rate (RCG), and the overall rate of crystallization (RCO). The crystal nucleation rate is the growth rate of new crystals and the crystal growth rate is the rate of growth of formed crystals. The overall rate of crystallization is the sum of curves RN and RCG.
  • In certain embodiments, the temperature of the cooling polymer construct can be at a temperature at which the overall crystallization rate is relatively low. At such a temperature, the increase in crystallinity is predominantly due to formation of crystalline domains around the bioceramic particles, rather than the growth of existing crystals. In some embodiments, the temperature can be in a range in which the crystal nucleation rate is larger than the crystal growth rate. In one embodiment, the temperature can be in a range in which the crystal nucleation rate is substantially larger than the crystal growth rate. For example, the temperature can be where the ratio of the crystal nucleation rate to crystal growth rate is 2, 5, 10, 50, 100, or greater than 100. In another embodiment, the temperature range may be in range, ΔT, shown in FIG. 3, between about Tg to about 0.25(Tm−Tg)+Tg.
  • In general, good bonding between a continuous phase and a discrete or reinforcing phase in a composite material facilitates improvement of the mechanical performance of the composite. For example, increase of the modulus and fracture toughness of a polymer due to a bioceramic particle phase can be enhanced by good bonding between the polymer and particles.
  • In some embodiments, bioceramic particles may include an adhesion promoter to improve the adhesion between the particles and the polymer matrix. In one embodiment, an adhesion promoter can include a coupling agent. A coupling agent refers to a chemical substance capable of reacting with both the bioceramic particle and the polymer matrix of the composite material. A coupling agent acts as an interface between the polymer and the bioceramic particle to form a chemical bridge between the two to enhance adhesion.
  • The adhesion promoter may include, but is not limited to, silane and non-silane coupling agents. For example, the adhesion promoter may include 3-aminopropyltrimethoxysilane, 3-aminopropyltriethoxysilane, aminopropylmethyldiethoxy silane, organotrialkoxysilanes, titanates, zirconates, and organic acid-chromium chloride coordination complexes. In particular, 3-aminopropyltrimethoxysilane has been shown to facilitate adhesion between poly(L-lactide) and bioglass. Biomaterials 25 (2004) 2489-2500.
  • In some embodiments, the surface of the bioceramic particles may be treated with an adhesion promoter prior to mixing with the polymer matrix. In one embodiment, the bioceramic particles can be treated with a solution containing the adhesion promoter. Treating can include, but is not limited to, coating, dipping, or spraying the particles with an adhesion promoter or a solution including the adhesion promoter. The particles can also be treated with a gas containing the adhesion promoter. In one embodiment, treatment of the bioceramic particles includes mixing the adhesion promoter with solution of distilled water and a solvent such as ethanol and then adding bioceramic particles. The bioceramic particles can then be separated from the solution, for example, by a centrifuge, and the particles can be dried. The bioceramic particles may then used to form a polymer composite. In an alternative embodiment, the adhesion promoter can be added to the particles during formation of the composite. For example, the adhesion promoter can be mixed with a bioceramic/polymer mixture during extrusion.
  • As indicated above, a device may be composed in whole or in part of materials that degrade, erode, or disintegrate through exposure to physiological conditions within the body until the treatment regimen is completed. The device may be configured to disintegrate and disappear from the region of implantation once treatment is completed. The device may disintegrate by one or more mechanisms including, but not limited to, dissolution and chemical breakdown.
  • The duration of a treatment period depends on the bodily disorder that is being treated. For illustrative purposes only, in treatment of coronary heart disease involving use of stents in diseased vessels, the duration can be in a range from about a month to a few years. However, the duration is typically in a range from about six to twelve months. Thus, it is desirable for an implantable medical device, such as a stent, to have a degradation time at or near the duration of treatment. Degradation time refers to the time for an implantable medical device to substantially or completely erode away from an implant site.
  • Several mechanisms may be relied upon for erosion and disintegration of implantable devices which include, but are not limited to, mechanical, chemical breakdown and dissolution. Therefore, bodily conditions can include, but are not limited to, all conditions associated with bodily fluids (contact with fluids, flow of fluids) and mechanical forces arising from body tissue in direct and indirect contact with a device. Degradation of polymeric materials principally involves chemical breakdown involving enzymatic and/or hydrolytic cleavage of device material due to exposure to bodily fluids such as blood.
  • Chemical breakdown of biodegradable polymers results in changes of physical and chemical properties of the polymer, for example, following exposure to bodily fluids in a vascular environment. Chemical breakdown may be caused by, for example, hydrolysis and/or metabolic processes. Hydrolysis is a chemical process in which a molecule is cleaved into two parts by the addition of a molecule of water. Consequently, the degree of degradation in the bulk of a polymer is strongly dependent on the diffusivity, and hence the diffusion rate of water in the polymer.
  • Another deficiency of some biodegradable polymers, such as poly(L-lactide), is that the degradation rate is slow and results in a degradation time of a stent outside of the desired range. A preferred degradation is from six to twelve months. Increasing the equilibrium content of moisture in a biodegradable polymer that degrades by hydrolysis can increase the degradation rate of a polymer. Various embodiments of the present invention include increasing the equilibrium moisture content in a polymer of a device to accelerate the degradation rate.
  • In some embodiments, bioabsorbable bioceramic particles may be included in a bioceramic/polymer composite device to increase the degradation rate of the polymer and to decrease the degradation time of a device made from the composite. In an embodiment, the degradation rate of a bioceramic/polymer composite device can be tuned and/or adjusted to a desired time frame. As the bioceramic particle erodes within the polymeric matrix, the porosity of the matrix increases. The increased porosity increases the diffusion rate of moisture through the polymeric matrix, and thus, the equilibrium moisture content of the polymeric matrix. As a result, the degradation rate of the polymer is increased. The porous structure also increases the transport of degradation products out of the matrix, which also increases the degradation rate of the matrix.
  • In certain embodiments, the degradation rate and degradation time of the device can be tuned or controlled through variables such as the type of bioceramic material and the size and shape of particles. In some embodiments, bioceramic materials can be selected to have a higher degradation rate than the polymer matrix. The faster the degradation rate of the bioceramic material, the faster the porosity of the polymer matrix increases which results in a greater increase in the degradation rate of the polymer matrix. Additionally, the size of the particles influence the time for erosion of the particles. The smaller the particles, the faster the erosion of the particles because of the higher surface area per unit mass of particles.
  • For example, nanoparticles may have a relatively fast erosion rate compared to microparticles. Additionally, elongated particles, such as fibers, may tend to erode faster on a per unit mass basis due to the higher surface area per unit mass of the particle. Also, short fibers may tend to erode faster than longer fibers. Short fibers refer to long fibers than have been cut into short lengths. In some embodiments, the short fibers may be made by forming fibers as described above, and cutting them into short lengths. In one embodiment, a length of at least a portion of the short fibers is substantially smaller than a diameter of the formed tube. For example, in some embodiments, the short fibers may be less than 0.05 mm long. In other embodiments, the short fibers may be between 0.05 and 8 mm long, or more narrowly between 0.1 and 0.4 mm long or 0.3 and 0.4 mm long.
  • Furthermore, the size and distribution of pores created by erosion of bioceramic particles can also influence the degradation rate and time of the polymer matrix. Smaller particles, such as nanoparticles, create a porous network that exposes a larger volume of polymer matrix to bodily fluid than larger particles, like microparticles. As a result the degradation rate and time of the matrix may be higher when nanoparticles are used rather than microparticles.
  • Through appropriate selection of the type of material for the particles and the size and shape of the particles, the particles and the device can be designed to have a selected erosion rates and degradation time. For example, the particles can designed erode away in several minutes, hours, days, or a month upon exposure to bodily fluid.
  • As indicated above, many biodegradable polymers degrade by the mechanism of hydrolysis. The rate of the hydrolysis reaction tends to increase as the pH decreases. Since the degradation products of such polymers as polylactides are acidic, the degradation products have an autocatalytic effect. Therefore, the pH of the degradation products of the bioceramics can also affect the degradation rate of a device. Therefore, bioceramic particles with acidic degradation by-products may further increase the rate of degradation of a matrix polymer.
  • For example, tricalcium phosphate releases acidic degradation products. Thus, some embodiments may include a composite including a bioceramic having acidic degradation products upon exposure to bodily fluids. The acidic degradation products can increase the degradation rate of the polymer which can decrease the degradation time of the device.
  • In other embodiments, a composite can have bioceramic particles that have basic degradation products. For example, hydroxyapatite releases basic degradation products. The basic degradation products of the bioceramic particles can reduce the autocatalytic effect of the polymer degradation by neutralizing the acidic degradation products of the polymer degradation. In some embodiments, the basic degradation products of the bioceramic particles can reduce the degradation rate of the polymer. Additionally, bioceramic particles having a basic degradation product may also depress foreign body reaction.
  • For example, in rapidly eroding implantable medical devices, such as, for example poly(lactide-co-glycolide) which can potentially produce a local pH drop due to the rapid release of acidic degradation products, the use of bioceramic particles having a basic degradation product may buffer the reaction and neutralize the local pH drop.
  • Furthermore, some semi-crystalline biodegradable polymers have a degradation rate that is slower than desired for certain stent treatments. As a result, the degradation time of a stent made from such polymers can be longer than desired. For example, a stent made from poly(L-lactide) can have a degradation time of between about two and three years. In some treatment situations, a degradation time of less than a year may be desirable, for example, between four and eight months.
  • As discussed above, the degradation of a hydrolytically degradable polymer follows a sequence including water penetration into the polymer followed by hydrolysis of bonds in the polymer. Thus, the degradation of a polymer can be influenced by its affinity for water and the diffusion rate of water through the polymer. A hydrophobic polymer has a low affinity for water which results in a relatively low water penetration. In addition, the diffusion rate of water through crystalline regions of a polymer is lower than amorphous regions. Thus, as either the affinity of a polymer for water decreases or the crystallinity increases, water penetration and water content of a polymer decreases.
  • Further embodiments of a biodegradable implantable medical device may be fabricated from a copolymer. In certain embodiments, the copolymer can be a matrix in a bioceramic/polymer composite. The copolymer can include hydrolytically degradable monomers or functional groups that provide desired degradation characteristics. For instance, the copolymer can include functional groups that increase water penetration and water content of the copolymer. In some embodiments, a copolymer can include a primary functional group and at least one additional secondary functional group. In one embodiment, the copolymer may be a random copolymer including the primary functional group and at least one additional secondary functional group.
  • In an embodiment, the copolymer with at least one secondary functional group can have a higher degradation rate than a homopolymer composed of the primary functional group. A stent fabricated from the copolymer can have a lower degradation time than a stent fabricated from a homopolymer composed of the primary functional group.
  • In an embodiment, the weight percent of the secondary functional group can be selected or adjusted to obtain a desired degradation rate of the copolymer or degradation time of a stent made from the copolymer. In some exemplary embodiments, the weight percent of the secondary functional group in a copolymer can be at least 1%, 5%, 10%, 15%, 30%, 40%, or, at least 50%. In certain exemplary embodiments, a secondary functional group can be selected and the weight percent of the a secondary functional group can be adjusted so that the degradation time of a stent, with or without dispersed bioceramic particles, can be less than 18 months, 12 months, 8 months, 5 months, or more narrowly, less than 3 month.
  • In some embodiments, the copolymer with at least one secondary functional group can have a lower crystallinity than a homopolymer composed of the primary functional group. It is believed that inclusion of a secondary functional group can perturb the crystalline structure of a polymer including the primary functional group, resulting in a reduced crystallinity. As a result, a stent fabricated from the copolymer has a larger percentage of amorphous regions, which allow greater water penetration. Thus, the degradation rate of the copolymer can be increased and the degradation time of a stent made from the copolymer can be decreased.
  • In one embodiment, a secondary functional group can be a stereoisomer of the primary functional group. One exemplary embodiment can include a poly(L-lactide-co-DL-lactide) copolymer. DL-lactide can be a secondary functional group that perturbs the crystalline structure of the poly(L-lactide) so that the copolymer has a lower crystallinity than the poly(L-lactide).
  • In some embodiments, increasing the number of secondary functional groups in the copolymer can result in a decrease in modulus of the copolymer as compared to a homopolymer of the primary functional group. The decrease in modulus can be due to the decrease in crystallinity. The decrease in modulus can reduce the ability of a stent to support a vessel. Thus, the weight percent of the secondary functional group can be adjusted so that the ability of a stent to act as structural support is not substantially reduced. The inclusion of bioceramic particles in the copolymer can partially or completely compensate for the reduction in the modulus of the copolymer.
  • In other embodiments, the copolymer can include at least one secondary functional group with a greater affinity for water than the primary functional group. The secondary functional group can be less hydrophobic or more hydrophilic than the primary functional group. The decreased hydrophobicity or increased hydrophilicity can increase the concentration of water near bonds prone to hydrolysis, increasing the degradation rate and lowering the degradation time of a stent made from the copolymer.
  • In certain embodiments, the secondary functional groups can be selected so that segments of the copolymer with a secondary functional group can degrade faster than the primary functional group segments. The difference in degradation rate can be due to the secondary functional groups being more hydrolytically active than the primary functional group. In one embodiment, a secondary functional group can be selected such that a homopolymer including the secondary functional group has a higher degradation rate than a homopolymer including the primary functional group.
  • In an exemplary embodiment, the copolymer can be poly(L-lactide-co-glycolide). The primary functional group can be L-lactide and the secondary functional group can be glycolide. The weight percent of the glycolide in the copolymer can be at least 1%, 5%, 10%, 15%, 30%, 40%, or, at least 50%. In certain exemplary embodiments, the weight percent of glycolide group can be adjusted so that the degradation time of a stent, with or without dispersed bioceramic particles, can be less than 18 months, 12 months, 8 months, 5 months, or more narrowly, 3 months or less.
  • Further embodiments of the invention include formation of a bioceramic/polymer composite and fabrication of an implantable medical device therefrom. As indicated above, a composite of a polymer and bioceramic particles can be extruded to form a polymer construct, such as a tube. A stent can then be fabricated from the tube. The composite can be formed in a number of ways. In some embodiments, the composite can be formed by melt blending. In melt blending the bioceramic particles are mixed with a polymer melt. The particles can be mixed with the polymer melt using extrusion or batch processing.
  • In one embodiment, the bioceramic particles can be combined with a polymer in a powdered or granular form prior to melting of the polymer. The particles and polymer can be mixed using mechanical mixing or stirring such as agitation of the particles and polymer in a container or a mixer. The agitated mixture can then be heated to a temperature above the melt temperature of the polymer in an extruder or using batch processing.
  • However, a problem with the mechanical mixing or stirring techniques is that the polymer and particles may be separated into separate regions or layers. This is particularly a problem with respect to smaller particles such as nanoparticles. Additionally, obtaining a uniform dispersion by mixing particles with a polymer melt as described, is that particles can agglomerate or form clusters. The agglomeration of bioceramic particles makes it difficult to disperse the particles within the composite. The presence of larger clusters in the composite tends to result in a decrease in material performance. Such larger clusters can result in the formation of voids in a composite device, which are preferential sites for crack initiation and failure. The mechanical mixing in a conventional single screw extruder or in batch processing can be insufficient to break up the clusters, resulting in a nonuniform mixture of bioceramic particles and polymer.
  • Various embodiments of forming a composite may be employed to increase the uniformity of dispersion of bioceramic particles within a polymer in a composite. One set of embodiments may include forming a composite from a suspension of bioceramic particles and a polymer solution. A composite formed using a suspension may result in a composite having more uniformly dispersed particles than the mixing methods described above.
  • Another set of embodiments can include reducing the agglomeration of bioceramic particles by decreasing the surface energy between particles to improve the dispersion of particles in the composite. In some embodiments, the surface energy can be decreased by treating the bioceramic particles with a surface modifier. In still another set of embodiments, agglomeration can be reduced through mechanical mixing that applies shear stress to the particles sufficient to reduce the size of clusters of particles.
  • With respect to the first set of embodiments, bioceramic particles can be mixed with a polymer by solution blending in which a composite mixture of bioceramic particles and polymer is formed from a suspension of particles in a polymer solution. Certain embodiments of a method of forming an implantable medical device may include forming a suspension including a fluid, a polymer, and bioceramic particles. A “suspension” is a mixture in which particles are suspended or dispersed in a fluid. The fluid can be a solvent for the polymer so that the polymer is dissolved in the fluid. The particles can be mixed with the fluid before or after dissolving the polymer in the fluid.
  • Various mechanical mixing methods known to those of skill in the art may be used to disperse the bioceramic particles in the suspension. In one embodiment, the suspension can be treated with ultrasound, for example, by an ultrasonic mixer. The method may further include removing some or all of the fluid or separating the polymer and bioceramic particles from the fluid to obtain a composite including bioceramic particles dispersed within the polymer.
  • In some embodiments, the polymer and bioceramic particles can be separated from the fluid by combining the suspension with a second fluid that may be a poor solvent for the polymer. At least some of the polymer may be allowed to precipitate upon combining the suspension solution with the second fluid. In some embodiments, at least some of the bioceramic particles may precipitate from the suspension with the precipitated polymer to form a composite mixture.
  • The precipitated composite mixture may then be filtered out of the solvents. The filtered composite mixture can be dried to remove residual solvents. For example, the composite mixture can be dried in a vacuum oven or by blowing heated gas on the mixture.
  • Exemplary polymers may include, but are not limited to, poly (L-lactic acid), poly (DL-lactic acid), poly(lactide-coglycolide). Representative solvents for such polymers can include toluene and chloroform. Representative poor solvents for these polymers that may be used to precipitate the polymer include methanol, ethanol, isopropanol, and various alkanes such as hexane or heptane.
  • It is believed that in a suspension including bioceramic nanoparticles, the particles can have strong interactions with polymer chains in solution which can result in particles becoming encapsulated or surrounded by polymer chains. Thus, when the polymer is precipitated from the solution, the interactions of the particles with the polymer can overcome interactions of the particles with the solution so that the particles precipitate with the polymer.
  • Additionally, it has been observed that the both the degree of precipitation of particles and the degree of dispersion of particles within the precipitated polymer depends upon the amount of polymer dissolved in the solution. The degree of precipitation refers to the amount of particles that precipitate out of the suspension. The degree of dispersion of particles within the precipitated polymer refers to the degree of mixing of the particles with the polymer.
  • The amount of polymer can be quantified by the weight percent of the polymer in the suspension solution. In addition, the viscosity of the solution is also related to the amount of polymer in the solution. The higher the weight percent of dissolved polymer, the higher is the viscosity of the suspension solution.
  • For a given concentration of suspended particles, as weight percent of dissolved polymer or viscosity is reduced, the degree of precipitation of particles is reduced. This is likely due to the reduced interaction of the particles with the polymer chains. Thus, at lower weight percent of polymer or viscosity, the amount of particles precipitating can be relatively low.
  • Additionally, for a given concentration of suspended particles, as the weight percent of polymer or viscosity of the solution is increased beyond an observed range, the degree of dispersion of particles in the precipitated polymer tends to decrease. It is believed that at higher weight percent of polymer or higher viscosity, the interactions between polymer chains reduce the interaction of particles with polymer chains that cause particles to precipitate. For example, particles may be unable to move freely among the polymer chains.
  • A given suspension can have a particular combination of type of particles, particle concentration, and solvent. For this given suspension, the polymer weight percent or viscosity that can be varied to obtain both a desired degree of precipitation of particles and degree of dispersion of particles in the precipitated polymer. Thus, there may be a range of polymer weight percent or viscosity that can result in a desired degree of precipitation of particles and degree of dispersion of particles in precipitated polymer.
  • Additionally, the manner of combining the suspension with the poor solvent can also affect the degree of precipitation and degree of dispersion. For example, depositing a fine mist of small droplets into a poor solvent can more readily result in a desired degree of precipitation and degree of dispersion. Thus, the manner of combining the suspension with the poor solvent can influence the range of polymer weight percent or viscosity that results in a desired degree of precipitation and degree of dispersion.
  • In other embodiments, the polymer and bioceramic particles can be separated from the fluid in the suspension by evaporating the fluid in the suspension solution. In some embodiments, at least some of the bioceramic particles may be separated from the suspension solution along with the polymer upon evaporation to form the composite mixture. In one embodiment, the solvent can be removed by evaporation at high vacuum in a vacuum oven.
  • In further embodiments, the composite formed from the suspension solution can be conveyed into an extruder. The composite mixture may be extruded at a temperature above the melting temperature of the polymer and less than the melting temperature of the bioceramic particles. In some embodiments, the dried composite mixture may be broken into small pieces by, for example, chopping or grinding. Extruding smaller pieces of the composite mixture may lead to a more uniform distribution of the nanoparticles during the extrusion process.
  • The extruded composite mixture may then be formed into a polymer construct, such as a tube or sheet which can be rolled or bonded to form a tube. A medical device may then be fabricated from the construct. For example, a stent can be fabricated from a tube by laser machining a pattern in to the tube.
  • In another embodiment, a polymer construct may be formed from the composite mixture using an injection molding apparatus.
  • Preparation of a desired amount of precipitated composite mixture may require a large amount of solvent and precipitant. Therefore, in some embodiments, it may be advantageous to melt blend precipitated composite mixture with an amount of polymer in an extruder or in a batch process. The polymer can be the same or a different polymer of the precipitated composite mixture. For example, a relatively small amount of precipitated composite mixture that has a weight percent of bioceramic particles higher than is desired can be prepared. The precipitated composite mixture may be melt blended with an amount of biodegradable polymer to form a composite mixture than has a desired weight percent of bioceramic particles.
  • As indicated above, embodiments for increasing the dispersion of bioceramic particles in a composite can include decreasing the surface energy of particles with a surface modifier. Bioceramic particles tend to aggregate due to high surface energy resulting from strong interparticle interactions.
  • Certain embodiments can include processing a plurality of bioceramic particles with a non-reactive surface modifier to reduce the surface energy between the particles, hence, agglomeration of particles. A composite can then be formed with the processed bioceramic particles, the composite including the processed bioceramic particles dispersed within a polymer.
  • In some embodiments, the processing can include forming a suspension of bioceramic particles including a suitable solvent and a non-reactive surface modifier dissolved in the solvent. The non-reactive surface modifier tends to reduce surface energy between the particles, and thus, the fracture strength of the agglomerates or clusters of bioceramic particles. The non-reactive surface modifier may cause or facilitate the break-up or reduction in size of the agglomerates or clusters of particles.
  • Additionally, due to the decreased fracture strength, the agglomerates are more susceptible to break-up from application of conventional stirring techniques. In an embodiment, the breakdown of the agglomerates can be facilitated by mixing the suspension with an ultrasonic and/or high speed mixer known to one of skill in the art. The suspension can be mixed to obtain a desired degree of break-up of agglomerates. For example, the suspension can be mixed from one to five minutes, five to 30 minutes, 30 minutes to an hour, one hour to eight hours, or more than eight hours. The solvent may be removed and then mixed it with a polymer to form composite, e.g. by melt processing.
  • In one embodiment, the polymer of the composite can be dissolved in the solvent after forming the suspension with the non-reactive surface modifier and nano bioceramic. Alternatively, the polymer can be dissolved prior to adding the non-reactive surface modifier. After sufficient exposure of the bioceramic particles to the non-reactive surface modifier, the solvent can be removed to form a composite mixture of polymer and bioceramic particles. For example, the solvent can be removed as described above through precipitation or through evaporation.
  • Representative examples of non-reactive surface modifiers for treating bioceramic particles include, but are not limited to, stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide (PEO-b-PPO-b-PEO), and polyethylene oxide-b-polylactide. The non-reactive surface modifiers are used in amounts effective to prevent agglomeration of bioceramic particles. Generally, the non-reactive surface modifiers can be used in amounts of about 0.1 weight % to about 3 weight % by weight of the total composite mixture, more narrowly about 0.3 weight % to about 2 weight %, or more narrowly about 0.5 weight % to about 1 weight %.
  • As indicated above, further embodiments for decreasing agglomeration and increasing the dispersion of bioceramic particles in a composite can include processing a mixture of particles and polymer with mechanical methods sufficient to reduce agglomeration. Certain embodiments can include processing a mixture of a polymer and agglomerated bioceramic particles under high shear stress conditions. Some embodiments can include processing the mixture such that the particles are subjected to shear stress higher than the fracture strength of the agglomerated particles. In one embodiment, a polymer melt can blended or mixed with bioceramic particles in a manner that subjects the mixture to a shear stress higher than the fracture strength of agglomerates of bioceramic particles.
  • Generally, mixing a polymer melt with a melt viscosity η at a shear rate γ, subjects the melt and particles to a shear stress τ=η*γ. Thus, polymer/bioceramic mixture can be processed so that a maximum shear stress generated during the melt blending is higher than the fracture strength of the bioceramic particle agglomerates. Agglomerated particles may be mechanically broken down and more uniformly dispersed within the polymer.
  • As discussed above, the mechanical mixing of polymer and bioceramic particles provided by a single-screw extruder is generally insufficient to obtain a uniformly dispersed mixture of bioceramic particles with the polymer. The shear stress produced by a single screw extruder is typically lower than the fracture strength of bioceramic particle agglomerates.
  • Various kinds of mixing devices may be employed that can apply a shear stress higher than the fracture strength of agglomerates to a polymer melt. Mechanical blending devices that can apply a sufficiently high shear stress include, but are not limited to, a twin screw extruder or batch mixer. During blending, once the shear stress is higher than the fracture strength of bioceramic particles agglomerates, the agglomerates are broken down and more uniformly dispersed into the polymer. The polymer and bioceramic particles can be fed into a mechanical blending device separately and processed at high shear stress. Alternatively, a composite mixture of polymer and bioceramic particles can be fed into a mechanical blending device and processed at the high shear stress.
  • The polymer/bioceramic particle mixture can be processed at the sufficiently high shear stress for a time sufficient to reduce agglomeration and disperse the particles. For example, the mixture can be processed between about 5 min. to about 30 min., more narrowly about 8 min. to about 20 min., or more narrowly about 10 min to about 15 min.
  • In some embodiments, a composite of polymer and bioceramic particles formed by other methods disclosed herein can be processed at the high shear stress conditions. In one embodiment, the composite formed from a suspension solution through evaporation or precipitation can be processed at high shear stress conditions, for example, in a twin screw extruder or a batch mixer. In another embodiment, the composite formed with surface modified bioceramic particles can also be processed in this manner.
  • In some embodiments, bioceramic particles treated with a non-reactive surface modifier can be processed with a polymer to form a composite. The bioceramic particles can be treated by dispersing the bioceramic particles in a solution including the non-reactive surface modifier. Upon removal of the solvent from the solution, the treated particles can be processed in a mechanical mixing device such as an extruder. Since the fracture strength of clusters of agglomerated particles is decreased by the non-reactive surface modifier, even a conventional single screw extruder may facilitate break-up of the clusters, reducing agglomeration. In other embodiments, the treated particles and polymer can be processed in a device, such as a twin screw extruder or batch mixer, that subjects the particles and polymer to shear stress to shear stress greater than the fracture strength of clusters of untreated bioceramic particles.
  • Representative examples of polymers that may be used to fabricate an implantable medical device include, but are not limited to, poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate), poly(lactide-co-glycolide), poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride, poly(glycolic acid), poly(glycolide), poly(L-lactic acid), poly(L-lactide), poly(D,L-lactic acid), poly(L-lactide-co-glycolide); poly(D,L-lactide), poly(caprolactone), poly(trimethylene carbonate), polyethylene amide, polyethylene acrylate, poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin, fibrinogen, cellulose, starch, collagen and hyaluronic acid), polyurethanes, silicones, polyesters, polyolefins, polyisobutylene and ethylene-alphaolefin copolymers, acrylic polymers and copolymers other than polyacrylates, vinyl halide polymers and copolymers (such as polyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether), polyvinylidene halides (such as polyvinylidene chloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such as polystyrene), polyvinyl esters (such as polyvinyl acetate), acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon 66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides, polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, cellulose acetate, cellulose butyrate, cellulose acetate butyrate, cellophane, cellulose nitrate, cellulose propionate, cellulose ethers, and carboxymethyl cellulose.
  • Additional representative examples of polymers that may be especially well suited for use in fabricating an implantable medical device according to the methods disclosed herein include ethylene vinyl alcohol copolymer (commonly known by the generic name EVOH or by the trade name EVAL), poly(butyl methacrylate), poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508, available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride (otherwise known as KYNAR, available from ATOFINA Chemicals, Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethylene glycol.
  • The examples and experimental data set forth below are for illustrative purposes only and are in no way meant to limit the invention. The following examples are given to aid in understanding the invention, but it is to be understood that the invention is not limited to the particular materials or procedures of examples. In examples 1-7, hydroxyapatite nano particles (HAP) are used as the bioceramic nanoparticles. In examples 8-9 calcium sulfate nanoparticles are used as the nanoparticles. The polymers used in the Examples were poly(L-lactic acid) (PLLA), poly(DL-lactic acid) (PDLLA), and poly(lactic acid-co-glycolide) (PLGA).
  • EXAMPLE 1 Prophetic Example of Solution Blending of Polymer and Bioceramic Particles
  • Step 1: Add bioceramic particles into suitable solvent, such as chloroform, acetone, etc. and stir to form a bioceramic particle suspension solution.
  • Step 2: Slowly add a polymer such as PLLA, PDLLA, PLGA into suspension solution and stir until polymer dissolves completely. In this step, the solution may still have a relatively low viscosity. However, the bioceramic particles should be well dispersed while stirring.
  • Step 3: Slowly add the polymer into solution again to gradually increase solution viscosity. Repeat this step as needed until the polymer is completely dissolved and reasonable solution viscosity is developed.
  • Step 4: Apply ultrasonic mixing to suspension solution for 15-30 min to further disperse all the HAP uniformly into the PLLA solution.
  • Step 6: Add suspension solution to methanol to precipitate polymer and particles.
  • EXAMPLE 2
  • Solution Blending of PLLA/HAP Composite (100:1 wt/wt)
  • Step 1: Added 50 mg HAP particles into 300 mL of chloroform and stirred for 10-30 minutes to form bioceramic particle suspension solution.
  • Step 2: Slowly added 5 g PLLA into suspension solution and stirred about 8 hours to dissolve all polymer.
  • Step 3: Applied ultrasonic mixing to suspension solution for 15-30 min to further disperse HAP particles into PLLA solution.
  • Step 4: Added suspension solution to 1 L methanol to precipitate polymer and particles.
  • Step 5: Filtered the precipitate and dried about 24 hours in vacuum oven at 60° C. End product is PLLA/HAP composite.
  • Composites were also made with 2 wt % and 5 wt % HAP.
  • Mechanical Properties and Morphology of PLLA/HAP Composite (100:1 wt/wt)
  • Tensile testing of the composite samples and a pure PLLA were performed using an Instron tensile tester obtained from Instron in Canton, Mass. Test samples were prepared by hot pressing the PLLA/HAP composites and pure polymer to a thin film at 193° C. for 60 seconds. Testing bars were cut from the thin film and tested. The peak stress at break, the strain at break, and the Young's modulus were measured. The draw rate was about 0.5 in/min.
  • FIGS. 4-6 illustrate tensile testing results for pure PLLA and PLLA/HAP composites with 1 wt % HAP. FIG. 4 depicts the peak stress for the two samples. PLLA/HAP composites have a higher peak stress than the pure PLLA. FIG. 5 depicts the % strain at break for the two samples. The 1 wt % composite had a higher % strain at break than the pure PLLA. FIG. 6 depicts Young's modulus for the two samples. The 1 wt % samples had a higher Young's modulus than the pure PLLA.
  • EXAMPLE 3
  • Solution Blending of PLLA/HAP Composite (2:1 wt/wt) as HAP Intermedium Mixture
  • Step 1: Added 25 g HAP particles to 3 L chloroform and stirred for 10-30 minutes to form bioceramic particle suspension solution.
  • Step 2: Slowly added 50 g PLLA into suspension solution and stirred about 8 hours to dissolve all polymer.
  • Step 3: Applied ultrasonic mixing for 15-30 min to further disperse HAP particles into PLLA solution.
  • Step 4: Added suspension solution to 9 L methanol to precipitate particles and polymer.
  • Step 5: Filtered the precipitate and dried about 24 hours in vacuum oven at 60° C. End product is PLLA/HAP composite.
  • Extrusion of Precipitated PLLA/HAP (2:1 wt/wt) with PLLA
  • Step 1: Broke 2:1 wt/wt composite into small pieces
  • Step 2: Mixed 24 g of broken up composite and 376 g PLLA.
  • Step 3: Extruded mixture at 216° C.
  • EXAMPLE 4
  • Extrusion of Precipitated PLGA/HAP (2:1 wt/wt) with PLGA
  • Step 1: Broke PLGA/HAP composite into small pieces
  • Step 2: Mixed 12 g broken up PLGA/HAP composite and 396 g PLGA.
  • Step 3: Extruded mixture at 216° C.
  • EXAMPLE 5
  • As discussed above, to further improve mechanical properties of a bioceramic and polymer composite, the interfacial adhesion can be enhanced. The adhesion between bioceramic particles and a biodegradable polymer can be improved by coating at least a portion of the surfaces of the bioceramic particles with an adhesion promoter such as 3-aminopropyltrimethoxysilane and 3-aminopropyltriethoxysilane.
  • Example of the modification of HAP:
  • Step 1: Added 100 ml distillated water to 1900 ml Ethanol and stirred for 15-30 min.
  • Step 2: Added 20 g 3-aminopropyltrimethoxysilane to water-ethanol mixture and stirred for 1 h.
  • Step 3: Added 20 g HAP and stirred for 2 h.
  • Step 4: Centrifuged the modified HAP from solution.
  • Step 5: Dried HAP about 8 hours in vacuum oven.
  • EXAMPLE 6
  • A stent was fabricated from PLGA/nanoparticle composite tubing. Prior to cutting a stent pattern, the tubing was expanded at 109° C. in a blow molder to increase radial strength. A stent pattern was cut in the expanded tubing using an ultra-fast pulse laser. The stent was crimped at 30° C. After crimping, the stent was cold sterilized.
  • EXAMPLE 7
  • The mechanical properties of a stent fabricated from a PLGA/nanoparticle HAP composite were tested on an Instron compression tester. The recoil of stents was recorded after inflation and deflation of stent.
  • FIGS. 7 and 8 illustrate compression testing results for PLGA/HAP stent (100:1 wt/wt). The compression testing results for 100% PLGA is also included for comparison. FIG. 7 shows that the average radial strength of PLGA/HAP stent is about 11% higher than that of the 100% PLGA stent. FIG. 8 shows that the compression modulus of PLGA/nano HAP stent increased by about 23% over the 100% PLGA stent.
  • FIG. 9 shows the recoil of PLGA/nano HAP stent is about 15% less than the PLGA stent.
  • EXAMPLE 8
  • Stents were fabricated from a bioceramic/polymer composite with a matrix of PLGA copolymer. Calcium sulfate nanoparticles, which were pretreated by PEG-PPG-PEG surface modifier, were mixed with the copolymer matrix. PEG refers to polyethylene glycol and PPG refers to polypropylene glycol. The copolymer had 85 wt % L-lactide and 15 wt % GA. The composite was fabricated according to methods described herein. The stents were fabricated from tubes made from the composite and the tubes was radially expanded. The expanded tubes were laser cut to form stents.
  • Stents were fabricated and tested having a ratio of copolymer to particles of 100:1 by weight. Five stent samples were used for each testing (compression, recoil or expansion testing). at a zero time point (after fabrication) and after 16 hours of accelerated aging. Accelerated aging refers to aging at 40° C.
  • The zero time point samples were subjected to compression testing. Compression tests were performed with an Instron testing machine obtained from Instron in Canton, Mass. A stent sample was placed between two flat plates. The plates were adjusted so that the distance between the plates was the diameter of the stent in an uncompressed state. The plates were then adjusted to compress the stent by 10%, 15%, 25%, and 50% of the uncompressed plate distance. A resistance force, the amount force in units of Newtons/mm that was necessary to keep the stent at each compression distance, was measured. The resistance force corresponds to a measure of the radial strength (Rs). The modulus was also measured. Table 1 provides the compression test results for the 100:1 samples at zero time point. The Rs and modulus correspond to 50% compression.
    TABLE 1
    Compression test results for 100:1 samples at zero time point.
    Stent # Rs (psi) Modulus (psi)
    1 5.867 365.5
    2 5.443 291.5
    3 6.027 316.9
    4 5.879 324.2
    5 5.539 310.1
    Avg 5.751 321.6
    Std dev 0.248 27.360
  • The stent samples were subjected to recoil testing at both time points for each bioceramic composition. Each stent sample was deployed inside a length of Tecoflex elastic tubing. Tecoflex tubing can be obtained from Noveon, Inc., Cleveland, Ohio. The inside diameter (ID) of the Tecoflex tubing was 0.118 in and the outside diameter (OD) was 0.134 in. The ID of the sheath to hold the stent after crimping was 0.053 in. The stents were deployed to an outer diameter of 3.0 mm by inflating the balloon. The balloon was deflated, allowing the stent to recoil. The diameter of the recoiled stent was then measured.
  • The percent recoil was calculated from:
    % Recoil=(Inflated Diameter−Deflated Diameter)/Inflated Diameter×100%
  • The Inflated Diameter is the diameter of a deployed stent prior to deflating the balloon and the Deflated Diameter is the diameter of the recoiled stent. The deployment of the stent also results in an increase in length of the stent. The percent change in length was calculated from:
    % Length change=(Crimped length−Deployed length)/Crimped length×100%
  • Table 2 provides the results of the recoil test for the five 100:1 stent samples at zero time point. As shown in Table 1, the outside diameter (OD) of the stent when the balloon was inflated and deflated was measured at a proximal end, a middle, and distal end of the stent. ODwt refers to the outside diameter with tecoflex tubing and ODS refers to the outside diameter with stent only. The % Total Recoil given is the average of the recoil calculated at each point along the length of the stent.
    TABLE 2
    Recoil test results for 100:1 stent samples at zero time point.
    Crimp Inflated @ 8 atm Deflated Dimensions
    Tecoflex Tecoflex Stent Stent Prox Mid Distal Prox Mid Distal % %
    DWT DWT L L OD OD OD Stent L OD OD OD Total Length
    Stent # Prox Dist (mm) (mm) (mm) (mm) (mm) (mm) (mm) (mm) (mm) Recoil Change
    1 0.375 0.370 12.796 12.384 3.690 3.774 3.802 12.392 3.448 3.504 3.529 7.7% 3.2%
    2 0.368 0.378 12.833 12.529 3.792 3.785 3.811 12.450 3.541 3.487 3.496 8.4% 3.0%
    3 0.368 0.374 12.817 12.553 3.835 3.886 3.871 12.720 3.618 3.561 3.519 8.5% 0.8%
    4 0.375 0.370 12.863 12.492 3.792 3.806 3.871 12.774 3.653 3.673 3.644 4.8% 0.7%
    5 0.365 0.370 12.793 12.421 3.803 3.815 3.856 12.407 3.579 3.576 3.640 6.5% 3.0%
    ODwT 3.813 3.565 AVG 7.2% 2.1%
    ODS 3.441 3.193 Std 1.6% 1.3%
    Dev
  • The stents were subject to expansion testing at both time points. Each stent sample was deployed to an outer diameter of first 3.0 mm, then 3.5 mm, and then 4.0 mm. The number of cracks in the stents were counted at each deployment diameter.
  • Table 3 provides the number of cracks observed in the 100:1 stent samples at zero time point when deployed at a diameter of 3.5 mm. The columns refer to the size of the cracks: “Micro” refers to micro-sized cracks, “<25%” refers to cracks less than 25% of the strut width, “<50%” refers to cracks less than 50% of the strut width, and “>50%” refers to cracks greater than 50% of the strut width. “A” refers to the “v-shaped” region or bending region of struts and “B” refers to a spider region at the intersection of 5 struts. The regions can be seen in FIGS. 10A-D. No broken struts were observed at either 3.5 mm or 4.0 mm deployment diameters. FIGS. 10A-D depict photographic images of the 100:1 stent sample 1 at zero time point before expansion, expanded to 3.5 mm, after recoil, and after expansion to 4.0 mm, respectively.
    TABLE 3
    Number of cracks observed at 3.5 mm deployment
    for 100:1 stent samples at zero time point.
    Stents Deployed to 3.5 mm
    Stent # Micro <25% <50% >50%
    1
    A 4 2
    B
    2
    A 2
    B
    3
    A 4
    B
    4
    A
    B
    5
    A 1
    B
  • Test Results for 100:1 Stent Samples at 16 Hours of Accelerated Aging
  • Table 4 provides the results of the recoil test for the 100:1 stent samples at 16 hours of accelerated aging. Tables 5A-B provide the number of cracks observed in the 100:1 stent samples at 16 hours of accelerated aging when deployed at a diameter of 3.5 mm and 4.0 mm, respectively. No broken struts were observed at either deployment diameter. FIGS. 11A-D depict photographic images of the 100:1 stent sample 1 at 16 hours of accelerated aging before expansion, expanded to 3.5 mm, after recoil, and after expansion to 4.0 mm, respectively.
    TABLE 4
    Recoil test results for 100:1 stent samples at 16 hours of accelerated aging.
    Crimp Dimensions Inflated @ 8 atm Dimensions Deflated Dimensions
    Tecoflex Tecoflex Stent Stent Prox Mid Distal Prox Mid Distal % %
    DWT DWT L L OD OD OD Stent L OD OD OD Total Length
    Stent # Prox Dist (mm) (mm) (mm) (mm) (mm) (mm) (mm) (mm) (mm) Recoil Change
    1 0.372 0.377 12.683 12.133 3.900 3.834 3.836 12.311 3.508 3.520 3.522 9.8% 2.9%
    2 0.380 0.383 12.830 12.306 3.870 3.809 3.780 12.380 3.623 3.546 3.584 6.8% 3.5%
    3 0.374 0.372 12.924 12.574 3.845 3.804 3.783 12.620 3.477 3.448 3.422 10.5% 2.4%
    4 0.373 0.375 12.806 12.411 3.933 3.874 3.881 12.486 3.526 3.508 3.508 10.8% 2.5%
    5 0.381 0.382 12.745 12.358 3.853 3.799 3.809 12.570 3.482 3.489 3.471 9.9% 1.4%
    ODwT 3.841 3.509 AVG 9.6% 2.5%
    ODS 3.464 3.132 Std Dev 1.6% 0.8%
  • TABLE 5A
    Number of cracks observed for 100:1 stent samples at 16
    hours of accelerated aging deployed at 3.5 mm.
    Stents Deployed to 3.5 mm
    Stent # Micro <25% <50% >50%
    1
    A 2 1
    B 1
    2
    A
    B
    3
    A 2 1
    B
    4
    A 1
    B
    5
    A 7
    B 1
  • TABLE 5B
    Number of cracks observed for stents deployed at 4.0 mm for
    100:1 stent samples at 16 hours of accelerated aging.
    Stents Deployed to 4.0 mm
    Stent # Micro <25% <50% >50%
    1
    A 6 1
    B 3 1
    2
    A
    B 1
    3
    A 5 1
    B 1
    4
    A 1 1
    B 2 1
    5
    A 4 2
    B 1
  • EXAMPLE 9
  • A stent was fabricated from a polymer/bioceramic composite. The matrix was a PLGA copolymer. Calcium sulfate nanoparticles were mixed with the copolymer matrix. The copolymer was 50 wt % L-lactide and 50 wt % glycolide. The ratio of copolymer to particles was 100:3 by weight. The composite was fabricated according to methods described herein. A tube was fabricated from the composite and the tube was radially expanded. The expanded tube was laser cut to form a stent. The expected degradation time for the composite stent is about 3-5 months.
  • EXAMPLE 10
  • The following example illustrates formation of a polymer/bioceramic particle composite with surface modified bioceramic particles.
  • Step 1: A solution was formed by dissolving 8 g stearic acid (surface modifier) in 200 ml isopropanol (solvent).
  • Step 2: 20 g Calcium sulfate (CS) nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3: The suspension was subjected to ultrasonic mixing for 30-60 min to reduce the agglomeration of nanoparticles.
  • Step 4: The solvent was removed through evaporation at high vacuum.
  • Step 5: A polymer/nanoparticle composite was prepared by blending two portions of the modified calcium sulfate nanoparticles and 100 portions of a PLLA composite mixture which contained uniformly dispersed nanoparticles in the PLLA matrix.
  • EXAMPLE 11
  • Preparation of PEO-b-PPO-b-PEO Modified Calcium Sulfate Master Batch Material
  • Step 1: A solution was formed by dissolving 20 g PEO-b-PPO-b-PPO (surface modifier) in 400 ml chloroform (solvent).
  • Step 2: 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3: The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4: A polymer/nanoparticle composite was prepared by adding 20 g LPLG to the modified calcium sulfate nanoparticles. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 5: The solvent was removed through evaporation at high vacuum.
  • EXAMPLE 12
  • LPLG/Modified Calcium Sulfate Tubing Preparation and Morphology Observation
  • Step 1: A solution was formed by dissolving 20 g PEO-b-PPO-b-PPO (surface modifier) in 400 ml chloroform (solvent).
  • Step 2: 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3: The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4: The solvent was removed through evaporation at high vacuum.
  • Step 5: A polymer/nanoparticle composite was prepared by mixing the modified calcium sulfate nanoparticles with 1000 g LPLG. The tubing was prepared through extrusion using single screw extruder.
  • FIG. 12A shows s SEM photographic image of an LPLG/unmodified CS particle composite. The CS particles size range from 1 to 50 μm, with most particles between 2 to 30 μm. FIG. 12B illustrates a photographic image of an LPLG/modified CS particle composite with particles modified with PEO-b-PPO-b-PEO. The CS particles size range from 0.1 to 10 μm, where most of the particles are from 0.3 to 6 μm.
  • EXAMPLE 13
  • Preparation of Stearic Acid Modified Calcium Sulfate Master Batch Material
  • Step 1: A solution was formed by dissolving 16 g stearic acid (surface modifier) in 400 ml chloroform.
  • Step 2: 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3: The suspension was subjected to ultrasonic mixing for 30 min.
  • Step 5: A polymer/nanoparticle composite was prepared by adding 20 g LPLG to the modified calcium sulfate nanoparticles. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 4: The solvent was removed through evaporation at high vacuum.
  • EXAMPLE 14
  • LPLG/Modified Calcium Sulfate Tubing Preparation and Morphology Observation
  • Step 1: A solution was formed by dissolving 16 g stearic acid (surface modifier) in 400 ml chloroform (solvent).
  • Step 2: 20 g Calcium sulfate nanoparticles were added to the solution to form a suspension. The suspension was stirred for 1 h.
  • Step 3: The suspension was subjected to ultrasonic mixing for 30 min to reduce the agglomeration of nanoparticles.
  • Step 4: The solvent was removed through evaporation at high vacuum.
  • Step 5: A polymer/nanoparticle composite was prepared by mixing the modified calcium sulfate nanoparticles with 1000 g LPLG. The reaction mixture was stirred overnight till all LPLG dissolved into solution.
  • Step 6: The cross-section morphology of the tubing was observed by SEM (FIG. 1C).
  • FIG. 12C shows a photographic image of LPLG/stearic acid modified CS particle composite, which is similar to the morphology of LPLG/PEO-b-PPO-b-PEO modified CS particle composite. The CS particles sizes range from 0.1 to 10 μm, where most of them are from 0.5 to 6 μm.
  • EXAMPLE 15
  • LPLG material and surface modified or non-modified CS nano particles were mixed and extruded through a twin screw extruder. CS content in the extruded material is roughly 2%. The morphology of the extruded material was observed by SEM
  • FIG. 13A illustrates a photographic image of extruded LPLG/CS composite using a twin screw extruder. Overall, the CS particles sizes range from 0.2 to 20 μm with most of the particles from 0.5 to 5 μm. As a comparison, FIG. 13B shows a photographic image of LPLG/CS particle composite using a single screw extruder with low shear stress. Overall, the CS particles sizes range from 1 to 50 μm, where most of them are from 2 to 30 μm. FIG. 13C shows an image of an LPLG/PEO-b-PPO-b-PEO modified CS particles composite by twin screw extruder at with high shear stress, the CS particles sizes range from 0.1 to 10 μm, where most of them are from 0.2 to 3 μm.
  • While particular embodiments of the present invention have been shown and described, it will be obvious to those skilled in the art that changes and modifications can be made without departing from this invention in its broader aspects. Therefore, the appended claims are to encompass within their scope all such changes and modifications as fall within the true spirit and scope of this invention.

Claims (47)

1. A method of fabricating an implantable medical device comprising:
processing a plurality of agglomerated bioceramic particles with a non-reactive surface modifier, the non-reactive surface modifier reducing the agglomeration;
forming a composite with the processed bioceramic particles, the composite including the processed bioceramic particles dispersed within a polymer; and
fabricating an implantable medical device from the composite.
2. The method according to claim 1, wherein the non-reactive surface modifier lowers the surface energy of the particles.
3. The method according to claim 1, wherein the processing comprises dispersing the bioceramic particles in a solution including the non-reactive surface modifier.
4. The method according to claim 1, wherein the processing comprises mixing a suspension having the non-reactive surface modifier and the plurality of bioceramic particles so that the agglomeration is reduced.
5. The method according to claim 1, wherein forming the composite comprises processing a mixture of the polymer and the processed particles with a shear stress higher than the fracture strength of clusters of agglomerated bioceramic particles so that agglomeration is further reduced.
6. The method according to claim 1, wherein forming the composite comprises processing a mixture of the polymer and the processed particles with a twin-screw extruder or a kneader in such a way that agglomeration is further reduced.
7. The method according to claim 1, wherein the non-reactive surface modifier is selected from the group consisting of stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide, and polyethylene oxide-b-polylactide.
8. The method according to claim 1, wherein the medical device is a stent.
9. The method according to claim 1, wherein the polymer is biodegradable.
10. The method according to claim 1, wherein the bioceramic particles are nanoparticles.
11. The method according to claim 1, wherein a ratio of the weight percent of the bioceramic particles to the polymer is at least about 1:200.
12. The method according to claim 1, wherein the bioceramic particles are selected from the group consisting of calcium sulfate and hydroxyapatite.
13. A method for fabricating an implantable medical device comprising:
forming a suspension solution including a fluid, a non-reactive surface modifier, a polymer and bioceramic particles, wherein the polymer is dissolved in the fluid, and wherein the bioceramic particles are dispersed in the solution;
removing all or substantially all of the fluid to form a composite mixture, wherein the composite mixture comprises the bioceramic particles dispersed within the polymer, wherein the non-reactive surface modifier reduces the agglomeration of the bioceramic particles in the suspension solution and/or during formation of the composite mixture; and
fabricating an implantable medical device from the composite mixture.
14. The method according to claim 13, wherein the fluid is removed by evaporating the fluid.
15. The method according to claim 13, wherein the non-reactive surface modifier lowers the surface energy of the particles.
16. The method according to claim 13, further comprising mixing the suspension solution having the non-reactive surface modifier and the plurality of bioceramic particles to facilitate reducing the agglomeration.
17. The method according to claim 13, further comprising processing the composite with a shear stress higher than the fracture strength of clusters of agglomerated bioceramic particles so that agglomeration is further reduced.
18. The method according to claim 13, further comprising processing the composite with a twin-screw extruder or a kneader in such a way that agglomeration is further reduced.
19. The method according to claim 13, wherein the non-reactive surface modifier is selected from the group consisting of stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide, and polyethylene oxide-b-polylactide.
20. The method according to claim 13, wherein the medical device is a stent.
21. The method according to claim 13, wherein the polymer is biodegradable.
22. The method according to claim 13, wherein the bioceramic particles are nanoparticles.
23. The method according to claim 13, wherein a ratio of the weight percent of the bioceramic particles to the polymer is at least about 1:200.
24. A stent comprising:
a structural element including a bioceramic/polymer composite, the composite having a plurality of bioceramic particles dispersed within a polymer, a non-reactive surface modifier being on a surface of at least some of the particles, wherein the surface modifier reduces agglomeration of the bioceramic particles prior to and during the formation of the composite.
25. The device according to claim 24, wherein the non-reactive surface modifier is selected from the group consisting of stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide, and polyethylene oxide-b-polylactide.
26. The device according to claim 24, wherein the medical device is a stent.
27. The device according to claim 24, wherein the polymer is biodegradable.
28. The device according to claim 24, wherein the bioceramic particles are nanoparticles.
29. The device according to claim 24, wherein a ratio of the weight percent of the bioceramic particles to the polymer is at least about 1:200.
30. A method of fabricating an implantable medical device comprising:
processing a mixture of a polymer and agglomerated bioceramic particles to form a composite of the bioceramic particles dispersed within the polymer, wherein the processing subjects the plurality of bioceramic particles to a shear stress higher than a fracture strength of clusters of the agglomerated bioceramic particles so that agglomeration is reduced; and
fabricating an implantable medical device from the composite.
31. The method according to claim 30, wherein the polymer and the agglomerated bioceramic particles are processed with a twin-screw extruder or a kneader.
32. The method according to claim 30, wherein the medical device is a stent.
33. The method according to claim 30, wherein the polymer is biodegradable.
34. The method according to claim 30, wherein the bioceramic particles are nanoparticles.
35. The method according to claim 30, wherein a ratio of the weight percent of the bioceramic particles to the polymer is at least about 1:200.
36. The method according to claim 30, wherein the bioceramic particles are selected from the group consisting of calcium sulfate and hydroxyapatite.
37. A method of fabricating an implantable medical device comprising:
processing a mixture of a polymer and bioceramic particles to form a composite, wherein the bioceramic particles have been treated with a non-reactive surface modifier that reduces the fracture strength of clusters of the bioceramic particles, the composite comprising the bioceramic particles dispersed within the polymer; and
fabricating an implantable medical device from the composite.
38. The method according to claim 37, wherein the non-reactive surface modifier lowers the surface energy of the particles.
39. The method according to claim 37, wherein the agglomerated bioceramic particles are treated by dispersing the bioceramic particles in a solution including the non-reactive surface modifier.
40. The method according to claim 37, wherein the processing subjects the composite to a shear stress higher than fracture strength of clusters of untreated bioceramic particles so that the agglomeration is reduced.
41. The method according to claim 37, wherein the processing subjects the composite to a shear stress higher than the fracture strength of clusters of the treated bioceramic particles so that the agglomeration is reduced.
42. The method according to claim 37, wherein the composite is processed in a kneader or a twin screw extruder.
43. The method according to claim 37, wherein the non-reactive surface modifier is selected from the group consisting of stearic acid, polyethylene oxide-b-polypropylene oxide-b-polyethylene oxide (PEO-b-PPO-b-PEO), and polyethylene oxide-b-polylactide.
44. The method according to claim 37, wherein the medical device is a stent.
45. The method according to claim 37, wherein the polymer is biodegradable.
46. The method according to claim 37, wherein the bioceramic particles are nanoparticles.
47. The method according to claim 43, wherein the bioceramic particles are selected from the group consisting of calcium sulfate and hydroxyapatite.
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US11/529,996 US7842737B2 (en) 2006-09-29 2006-09-29 Polymer blend-bioceramic composite implantable medical devices
US70655007A 2007-02-14 2007-02-14
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Cited By (23)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US20090088834A1 (en) * 2007-09-28 2009-04-02 Abbott Cardiovascular Systems Inc. Stent formed from bioerodible metal-bioceramic composite
US20090234432A1 (en) * 2008-03-12 2009-09-17 Pacetti Stephen D Polymer Stent With Break-Away Links For Enhanced Stent Retention
US20100168798A1 (en) * 2008-12-30 2010-07-01 Clineff Theodore D Bioactive composites of polymer and glass and method for making same
WO2012100651A1 (en) * 2011-01-27 2012-08-02 Dongguan Tiantianxiangshang Medical Technology Co., Ltd. Biodegradable stent formed with polymer-bioceramic nanoparticle composite and preparation method thereof
US20120277358A1 (en) * 2008-03-31 2012-11-01 Abbott Cardiovascular Systems Inc. Stent with nucleating agent
EP2397306A3 (en) * 2007-12-11 2013-03-27 Abbott Cardiovascular Systems Inc. Method of fabricating a stent from blow molded tubing
US8425591B1 (en) 2007-06-11 2013-04-23 Abbott Cardiovascular Systems Inc. Methods of forming polymer-bioceramic composite medical devices with bioceramic particles
US20140039603A1 (en) * 2011-05-13 2014-02-06 Abbott Cardiovascular Systems Inc. Bioabsorbable scaffolds made from composites
US8658082B2 (en) 2007-12-11 2014-02-25 Abbott Cardiovascular Systems Inc. Method of fabricating stents from blow molded tubing
US20150025619A1 (en) * 2007-01-19 2015-01-22 Elixir Medical Corporation Biodegradable endoprostheses and methods for their fabrication
US20150328373A1 (en) * 2014-05-19 2015-11-19 Abbott Cardiovascular Systems Inc. Additives To Increase Degradation Rate Of A Biodegradable Scaffolding And Methods Of Forming Same
CN105264014A (en) * 2013-05-16 2016-01-20 株式会社索夫塞拉 Biodegradable material
US9259339B1 (en) 2014-08-15 2016-02-16 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9381277B2 (en) 2009-04-23 2016-07-05 Purac Biochem Bv Resorbable and biocompatible fibre glass compositions and their uses
US9381275B2 (en) 2006-09-25 2016-07-05 Orthovita, Inc. Bioactive load-bearing composites
US9399086B2 (en) 2009-07-24 2016-07-26 Warsaw Orthopedic, Inc Implantable medical devices
US9480588B2 (en) 2014-08-15 2016-11-01 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9730819B2 (en) 2014-08-15 2017-08-15 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9855156B2 (en) 2014-08-15 2018-01-02 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9943426B2 (en) 2015-07-15 2018-04-17 Elixir Medical Corporation Uncaging stent
US10213949B2 (en) 2011-08-11 2019-02-26 Abbott Cardiovascular Systems Inc. Controlling moisture in and plasticization of bioresorbable polymer for melt processing
US10918505B2 (en) 2016-05-16 2021-02-16 Elixir Medical Corporation Uncaging stent
CN112545710A (en) * 2020-12-08 2021-03-26 中国人民解放军空军军医大学 Porous polyurethane support based on 3D prints

Families Citing this family (54)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US20080249633A1 (en) * 2006-08-22 2008-10-09 Tim Wu Biodegradable Materials and Methods of Use
US7731890B2 (en) 2006-06-15 2010-06-08 Advanced Cardiovascular Systems, Inc. Methods of fabricating stents with enhanced fracture toughness
US9517149B2 (en) 2004-07-26 2016-12-13 Abbott Cardiovascular Systems Inc. Biodegradable stent with enhanced fracture toughness
US8747879B2 (en) 2006-04-28 2014-06-10 Advanced Cardiovascular Systems, Inc. Method of fabricating an implantable medical device to reduce chance of late inflammatory response
US7971333B2 (en) 2006-05-30 2011-07-05 Advanced Cardiovascular Systems, Inc. Manufacturing process for polymetric stents
US20110066222A1 (en) * 2009-09-11 2011-03-17 Yunbing Wang Polymeric Stent and Method of Making Same
WO2007011708A2 (en) 2005-07-15 2007-01-25 Micell Technologies, Inc. Stent with polymer coating containing amorphous rapamycin
KR101406415B1 (en) 2005-07-15 2014-06-19 미셀 테크놀로지즈, 인코포레이티드 Polymer coatings containing drug powder of controlled morphology
WO2007127363A2 (en) 2006-04-26 2007-11-08 Micell Technologies, Inc. Coatings containing multiple drugs
US7959940B2 (en) * 2006-05-30 2011-06-14 Advanced Cardiovascular Systems, Inc. Polymer-bioceramic composite implantable medical devices
US8486135B2 (en) 2006-06-01 2013-07-16 Abbott Cardiovascular Systems Inc. Implantable medical devices fabricated from branched polymers
WO2008016670A2 (en) * 2006-08-01 2008-02-07 Abbott Cardiovascular Systems Inc. Implantable medical devices made from polymer-bioceramic composite
US9585989B2 (en) * 2006-09-19 2017-03-07 Boston Scientific Scimed, Inc. Ureteral stent having variable hardness
WO2008086369A1 (en) 2007-01-08 2008-07-17 Micell Technologies, Inc. Stents having biodegradable layers
US11426494B2 (en) 2007-01-08 2022-08-30 MT Acquisition Holdings LLC Stents having biodegradable layers
FR2912739B1 (en) * 2007-02-15 2012-10-12 Noraker PROCESS FOR PREPARING A COMPOSITE MATERIAL, MATERIAL OBTAINED AND APPLICATIONS
US7567129B2 (en) * 2007-06-29 2009-07-28 Intel Corporation Monolithic flexible power amplifier using integrated tunable matching networks
US7935143B2 (en) * 2008-01-02 2011-05-03 Abbott Cardiovascular Systems Inc. Stent formed from polymer-bioceramic composite with radiopaque bioceramic particles
US8685432B2 (en) * 2008-03-25 2014-04-01 University Of Utah Research Foundation Controlled release tissue graft combination biomaterials
JP5608160B2 (en) 2008-04-17 2014-10-15 ミセル テクノロジーズ、インコーポレイテッド Stent with bioabsorbable layer
CA2946195A1 (en) 2008-07-17 2010-01-21 Micell Technologies, Inc. Drug delivery medical device
CN102292053A (en) 2008-09-29 2011-12-21 卡迪尔克阀门技术公司 Heart valve
WO2010040009A1 (en) 2008-10-01 2010-04-08 Cardiaq Valve Technologies, Inc. Delivery system for vascular implant
EP2413847A4 (en) 2009-04-01 2013-11-27 Micell Technologies Inc Coated stents
EP4119098A1 (en) 2009-04-15 2023-01-18 Edwards Lifesciences CardiAQ LLC Vascular implant and delivery system
EP3366326A1 (en) 2009-04-17 2018-08-29 Micell Technologies, Inc. Stents having controlled elution
EP2453834A4 (en) 2009-07-16 2014-04-16 Micell Technologies Inc Drug delivery medical device
US11369498B2 (en) 2010-02-02 2022-06-28 MT Acquisition Holdings LLC Stent and stent delivery system with improved deliverability
CA2797110C (en) 2010-04-22 2020-07-21 Micell Technologies, Inc. Stents and other devices having extracellular matrix coating
US8579964B2 (en) 2010-05-05 2013-11-12 Neovasc Inc. Transcatheter mitral valve prosthesis
CA2805631C (en) 2010-07-16 2018-07-31 Micell Technologies, Inc. Drug delivery medical device
CN103477387B (en) 2011-02-14 2015-11-25 弗兰霍菲尔运输应用研究公司 Use the encoding scheme based on linear prediction of spectrum domain noise shaping
US9308087B2 (en) 2011-04-28 2016-04-12 Neovasc Tiara Inc. Sequentially deployed transcatheter mitral valve prosthesis
US9554897B2 (en) 2011-04-28 2017-01-31 Neovasc Tiara Inc. Methods and apparatus for engaging a valve prosthesis with tissue
WO2013003644A1 (en) * 2011-06-30 2013-01-03 Elixir Medical Corporation Biodegradable endoprostheses and methods for their fabrication
US10117972B2 (en) 2011-07-15 2018-11-06 Micell Technologies, Inc. Drug delivery medical device
US10188772B2 (en) 2011-10-18 2019-01-29 Micell Technologies, Inc. Drug delivery medical device
US9408952B2 (en) * 2011-11-30 2016-08-09 Abbott Cardiovascular Systems Inc. Pediatric application of bioabsorbable polymer stents in infants and children with congenital heart defects
CA2870188A1 (en) 2012-04-12 2013-10-17 Howard University Polylactide and calcium phosphate compositions and methods of making the same
US9345573B2 (en) 2012-05-30 2016-05-24 Neovasc Tiara Inc. Methods and apparatus for loading a prosthesis onto a delivery system
US9322295B2 (en) 2012-10-17 2016-04-26 General Electric Company Thermal energy storage unit with steam and gas turbine system
WO2014074854A1 (en) 2012-11-09 2014-05-15 Wang Tongxin Block copolymers for tooth enamel protection
US9889235B2 (en) 2013-02-05 2018-02-13 University Of Utah Research Foundation Implantable devices for bone or joint defects
US10583002B2 (en) 2013-03-11 2020-03-10 Neovasc Tiara Inc. Prosthetic valve with anti-pivoting mechanism
JP6330024B2 (en) 2013-03-12 2018-05-23 マイセル・テクノロジーズ,インコーポレイテッド Bioabsorbable biomedical implant
US9681951B2 (en) 2013-03-14 2017-06-20 Edwards Lifesciences Cardiaq Llc Prosthesis with outer skirt and anchors
US9572665B2 (en) 2013-04-04 2017-02-21 Neovasc Tiara Inc. Methods and apparatus for delivering a prosthetic valve to a beating heart
JP2016519965A (en) * 2013-05-15 2016-07-11 マイセル・テクノロジーズ,インコーポレイテッド Bioabsorbable biomedical implant
US11331191B2 (en) 2015-08-12 2022-05-17 Howmedica Osteonics Corp. Bioactive soft tissue implant and methods of manufacture and use thereof
CA2938576A1 (en) 2015-08-12 2017-02-12 Howmedica Osteonics Corp. Methods for forming scaffolds
EP3241571B1 (en) 2016-05-02 2020-07-22 Howmedica Osteonics Corporation Bioactive soft tissue implant and methods of manufacture and use thereof
JP6462822B2 (en) * 2017-10-12 2019-01-30 株式会社ソフセラ Biodegradable material
KR102172373B1 (en) * 2018-05-17 2020-11-02 서울대학교 산학협력단 A method for preparation of biodegradable implant with increasing its porosity over time
CN112441742A (en) * 2019-08-30 2021-03-05 江苏启灏医疗科技有限公司 Bioactive glass, nasal cavity stent composite material and application thereof

Citations (95)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4321711A (en) * 1978-10-18 1982-03-30 Sumitomo Electric Industries, Ltd. Vascular prosthesis
US4633873A (en) * 1984-04-26 1987-01-06 American Cyanamid Company Surgical repair mesh
US4656083A (en) * 1983-08-01 1987-04-07 Washington Research Foundation Plasma gas discharge treatment for improving the biocompatibility of biomaterials
US4718907A (en) * 1985-06-20 1988-01-12 Atrium Medical Corporation Vascular prosthesis having fluorinated coating with varying F/C ratio
US4722335A (en) * 1986-10-20 1988-02-02 Vilasi Joseph A Expandable endotracheal tube
US4723549A (en) * 1986-09-18 1988-02-09 Wholey Mark H Method and apparatus for dilating blood vessels
US4732152A (en) * 1984-12-05 1988-03-22 Medinvent S.A. Device for implantation and a method of implantation in a vessel using such device
US4733665A (en) * 1985-11-07 1988-03-29 Expandable Grafts Partnership Expandable intraluminal graft, and method and apparatus for implanting an expandable intraluminal graft
US4740207A (en) * 1986-09-10 1988-04-26 Kreamer Jeffry W Intralumenal graft
US4800882A (en) * 1987-03-13 1989-01-31 Cook Incorporated Endovascular stent and delivery system
US4816339A (en) * 1987-04-28 1989-03-28 Baxter International Inc. Multi-layered poly(tetrafluoroethylene)/elastomer materials useful for in vivo implantation
US4818559A (en) * 1985-08-08 1989-04-04 Sumitomo Chemical Company, Limited Method for producing endosseous implants
US4902289A (en) * 1982-04-19 1990-02-20 Massachusetts Institute Of Technology Multilayer bioreplaceable blood vessel prosthesis
US4994298A (en) * 1988-06-07 1991-02-19 Biogold Inc. Method of making a biocompatible prosthesis
US5084065A (en) * 1989-07-10 1992-01-28 Corvita Corporation Reinforced graft assembly
US5085629A (en) * 1988-10-06 1992-02-04 Medical Engineering Corporation Biodegradable stent
US5100429A (en) * 1989-04-28 1992-03-31 C. R. Bard, Inc. Endovascular stent and delivery system
US5104410A (en) * 1990-10-22 1992-04-14 Intermedics Orthopedics, Inc Surgical implant having multiple layers of sintered porous coating and method
US5108417A (en) * 1990-09-14 1992-04-28 Interface Biomedical Laboratories Corp. Anti-turbulent, anti-thrombogenic intravascular stent
US5108755A (en) * 1989-04-27 1992-04-28 Sri International Biodegradable composites for internal medical use
US5192311A (en) * 1988-04-25 1993-03-09 Angeion Corporation Medical implant and method of making
US5197977A (en) * 1984-01-30 1993-03-30 Meadox Medicals, Inc. Drug delivery collagen-impregnated synthetic vascular graft
US5279594A (en) * 1990-05-23 1994-01-18 Jackson Richard R Intubation devices with local anesthetic effect for medical use
US5282860A (en) * 1991-10-16 1994-02-01 Olympus Optical Co., Ltd. Stent tube for medical use
US5289831A (en) * 1989-03-09 1994-03-01 Vance Products Incorporated Surface-treated stent, catheter, cannula, and the like
US5290271A (en) * 1990-05-14 1994-03-01 Jernberg Gary R Surgical implant and method for controlled release of chemotherapeutic agents
US5306286A (en) * 1987-06-25 1994-04-26 Duke University Absorbable stent
US5306294A (en) * 1992-08-05 1994-04-26 Ultrasonic Sensing And Monitoring Systems, Inc. Stent construction of rolled configuration
US5383925A (en) * 1992-09-14 1995-01-24 Meadox Medicals, Inc. Three-dimensional braided soft tissue prosthesis
US5385580A (en) * 1990-08-28 1995-01-31 Meadox Medicals, Inc. Self-supporting woven vascular graft
US5389106A (en) * 1993-10-29 1995-02-14 Numed, Inc. Impermeable expandable intravascular stent
US5399666A (en) * 1994-04-21 1995-03-21 E. I. Du Pont De Nemours And Company Easily degradable star-block copolymers
US5407609A (en) * 1989-05-04 1995-04-18 Southern Research Institute Microencapsulation process and products therefrom
US5502158A (en) * 1988-08-08 1996-03-26 Ecopol, Llc Degradable polymer composition
US5591607A (en) * 1994-03-18 1997-01-07 Lynx Therapeutics, Inc. Oligonucleotide N3→P5' phosphoramidates: triplex DNA formation
US5591199A (en) * 1995-06-07 1997-01-07 Porter; Christopher H. Curable fiber composite stent and delivery system
US5593434A (en) * 1992-01-31 1997-01-14 Advanced Cardiovascular Systems, Inc. Stent capable of attachment within a body lumen
US5593403A (en) * 1994-09-14 1997-01-14 Scimed Life Systems Inc. Method for modifying a stent in an implanted site
US5599301A (en) * 1993-11-22 1997-02-04 Advanced Cardiovascular Systems, Inc. Motor control system for an automatic catheter inflation system
US5605696A (en) * 1995-03-30 1997-02-25 Advanced Cardiovascular Systems, Inc. Drug loaded polymeric material and method of manufacture
US5607442A (en) * 1995-11-13 1997-03-04 Isostent, Inc. Stent with improved radiopacity and appearance characteristics
US5607467A (en) * 1990-09-14 1997-03-04 Froix; Michael Expandable polymeric stent with memory and delivery apparatus and method
US5618299A (en) * 1993-04-23 1997-04-08 Advanced Cardiovascular Systems, Inc. Ratcheting stent
US5707385A (en) * 1994-11-16 1998-01-13 Advanced Cardiovascular Systems, Inc. Drug loaded elastic membrane and method for delivery
US5711763A (en) * 1991-02-20 1998-01-27 Tdk Corporation Composite biological implant of a ceramic material in a metal substrate
US5716981A (en) * 1993-07-19 1998-02-10 Angiogenesis Technologies, Inc. Anti-angiogenic compositions and methods of use
US5725549A (en) * 1994-03-11 1998-03-10 Advanced Cardiovascular Systems, Inc. Coiled stent with locking ends
US5726297A (en) * 1994-03-18 1998-03-10 Lynx Therapeutics, Inc. Oligodeoxyribonucleotide N3' P5' phosphoramidates
US5728751A (en) * 1996-11-25 1998-03-17 Meadox Medicals, Inc. Bonding bio-active materials to substrate surfaces
US5733326A (en) * 1996-05-28 1998-03-31 Cordis Corporation Composite material endoprosthesis
US5733925A (en) * 1993-01-28 1998-03-31 Neorx Corporation Therapeutic inhibitor of vascular smooth muscle cells
US5733330A (en) * 1997-01-13 1998-03-31 Advanced Cardiovascular Systems, Inc. Balloon-expandable, crush-resistant locking stent
US5733564A (en) * 1993-04-14 1998-03-31 Leiras Oy Method of treating endo-osteal materials with a bisphosphonate solution
US5741881A (en) * 1996-11-25 1998-04-21 Meadox Medicals, Inc. Process for preparing covalently bound-heparin containing polyurethane-peo-heparin coating compositions
US5855612A (en) * 1995-05-12 1999-01-05 Ohta Inc. Biocompatible titanium implant
US5855618A (en) * 1996-09-13 1999-01-05 Meadox Medicals, Inc. Polyurethanes grafted with polyethylene oxide chains containing covalently bonded heparin
US5858746A (en) * 1992-04-20 1999-01-12 Board Of Regents, The University Of Texas System Gels for encapsulation of biological materials
US5865814A (en) * 1995-06-07 1999-02-02 Medtronic, Inc. Blood contacting medical device and method
US5868781A (en) * 1996-10-22 1999-02-09 Scimed Life Systems, Inc. Locking stent
US5874109A (en) * 1994-07-27 1999-02-23 The Trustees Of The University Of Pennsylvania Incorporation of biological molecules into bioactive glasses
US5874165A (en) * 1996-06-03 1999-02-23 Gore Enterprise Holdings, Inc. Materials and method for the immobilization of bioactive species onto polymeric subtrates
US5873904A (en) * 1995-06-07 1999-02-23 Cook Incorporated Silver implantable medical device
US5874101A (en) * 1997-04-14 1999-02-23 Usbiomaterials Corp. Bioactive-gel compositions and methods
US5876743A (en) * 1995-03-21 1999-03-02 Den-Mat Corporation Biocompatible adhesion in tissue repair
US5877263A (en) * 1996-11-25 1999-03-02 Meadox Medicals, Inc. Process for preparing polymer coatings grafted with polyethylene oxide chains containing covalently bonded bio-active agents
US5879713A (en) * 1994-10-12 1999-03-09 Focal, Inc. Targeted delivery via biodegradable polymers
US5888533A (en) * 1995-10-27 1999-03-30 Atrix Laboratories, Inc. Non-polymeric sustained release delivery system
US5891192A (en) * 1997-05-22 1999-04-06 The Regents Of The University Of California Ion-implanted protein-coated intralumenal implants
US5897955A (en) * 1996-06-03 1999-04-27 Gore Hybrid Technologies, Inc. Materials and methods for the immobilization of bioactive species onto polymeric substrates
US6010445A (en) * 1997-09-11 2000-01-04 Implant Sciences Corporation Radioactive medical device and process
US6015541A (en) * 1997-11-03 2000-01-18 Micro Therapeutics, Inc. Radioactive embolizing compositions
US6042875A (en) * 1997-04-30 2000-03-28 Schneider (Usa) Inc. Drug-releasing coatings for medical devices
US6171609B1 (en) * 1995-02-15 2001-01-09 Neorx Corporation Therapeutic inhibitor of vascular smooth muscle cells
US6174330B1 (en) * 1997-08-01 2001-01-16 Schneider (Usa) Inc Bioabsorbable marker having radiopaque constituents
US6177523B1 (en) * 1999-07-14 2001-01-23 Cardiotech International, Inc. Functionalized polyurethanes
US6183505B1 (en) * 1999-03-11 2001-02-06 Medtronic Ave, Inc. Method of stent retention to a delivery catheter balloon-braided retainers
US6187045B1 (en) * 1999-02-10 2001-02-13 Thomas K. Fehring Enhanced biocompatible implants and alloys
US20020002399A1 (en) * 1999-12-22 2002-01-03 Huxel Shawn Thayer Removable stent for body lumens
US20020004101A1 (en) * 1995-04-19 2002-01-10 Schneider (Usa) Inc. Drug coating with topcoat
US20020004060A1 (en) * 1997-07-18 2002-01-10 Bernd Heublein Metallic implant which is degradable in vivo
US6511748B1 (en) * 1998-01-06 2003-01-28 Aderans Research Institute, Inc. Bioabsorbable fibers and reinforced composites produced therefrom
US6517888B1 (en) * 2000-11-28 2003-02-11 Scimed Life Systems, Inc. Method for manufacturing a medical device having a coated portion by laser ablation
US20030033001A1 (en) * 2001-02-27 2003-02-13 Keiji Igaki Stent holding member and stent feeding system
US6527801B1 (en) * 2000-04-13 2003-03-04 Advanced Cardiovascular Systems, Inc. Biodegradable drug delivery material for stent
US6537589B1 (en) * 2000-04-03 2003-03-25 Kyung Won Medical Co., Ltd. Calcium phosphate artificial bone as osteoconductive and biodegradable bone substitute material
US6676697B1 (en) * 1996-09-19 2004-01-13 Medinol Ltd. Stent with variable features to optimize support and method of making such stent
US6679980B1 (en) * 2001-06-13 2004-01-20 Advanced Cardiovascular Systems, Inc. Apparatus for electropolishing a stent
US6689375B1 (en) * 1999-11-09 2004-02-10 Coripharm Medizinprodukte Gmbh & Co. Kg Resorbable bone implant material and method for producing the same
US6689823B1 (en) * 1999-03-31 2004-02-10 The Brigham And Women's Hospital, Inc. Nanocomposite surgical materials and method of producing them
US6695920B1 (en) * 2001-06-27 2004-02-24 Advanced Cardiovascular Systems, Inc. Mandrel for supporting a stent and a method of using the mandrel to coat a stent
US6706273B1 (en) * 1999-08-14 2004-03-16 Ivoclar Vivadent Ag Composition for implantation into the human and animal body
US6709379B1 (en) * 1998-11-02 2004-03-23 Alcove Surfaces Gmbh Implant with cavities containing therapeutic agents
US6846323B2 (en) * 2003-05-15 2005-01-25 Advanced Cardiovascular Systems, Inc. Intravascular stent
US20070026069A1 (en) * 2003-03-28 2007-02-01 Shastri Venkatram P Biommetic hierarchies using functionalized nanoparticles as building blocks
US20080063685A1 (en) * 2006-09-13 2008-03-13 Yunbing Wang Degradable polymeric implantable medical devices with continuous phase and discrete phase

Family Cites Families (238)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US3414980A (en) * 1967-02-21 1968-12-10 Nat Dairy Prod Corp Method of spray drying
GB1237035A (en) 1969-08-20 1971-06-30 Tsi Travmatologii I Ortopedii Magnesium-base alloy for use in bone surgery
US3900632A (en) 1970-02-27 1975-08-19 Kimberly Clark Co Laminate of tissue and random laid continuous filament web
DE2010115A1 (en) 1970-03-04 1971-09-16 Farbenfabriken Bayer Ag, 5090 Leverkusen Process for the production of micro-granules
US3839743A (en) 1972-04-21 1974-10-08 A Schwarcz Method for maintaining the normal integrity of blood
US4104410A (en) 1973-12-21 1978-08-01 Malecki George J Processing of green vegetables for color retention in canning
US4110497A (en) 1976-07-02 1978-08-29 Snyder Manufacturing Co., Ltd. Striped laminate and method and apparatus for making same
DE2928007A1 (en) 1979-07-11 1981-01-15 Riess Guido Dr BONE IMPLANT BODY FOR PROSTHESES AND BONE CONNECTORS AND METHOD FOR THE PRODUCTION THEREOF
US4346028A (en) 1979-12-14 1982-08-24 Monsanto Company Asbestiform crystalline calcium sodium or lithium phosphate, preparation and compositions
DE3019996A1 (en) 1980-05-24 1981-12-03 Institute für Textil- und Faserforschung Stuttgart, 7410 Reutlingen HOHLORGAN
US4517687A (en) 1982-09-15 1985-05-21 Meadox Medicals, Inc. Synthetic woven double-velour graft
US4594407A (en) 1983-09-20 1986-06-10 Allied Corporation Prosthetic devices derived from krebs-cycle dicarboxylic acids and diols
US4596574A (en) 1984-05-14 1986-06-24 The Regents Of The University Of California Biodegradable porous ceramic delivery system for bone morphogenetic protein
CH671337A5 (en) 1984-06-19 1989-08-31 Ceskoslovenska Akademie Ved
US4879135A (en) 1984-07-23 1989-11-07 University Of Medicine And Dentistry Of New Jersey Drug bonded prosthesis and process for producing same
US4743252A (en) 1986-01-13 1988-05-10 Corvita Corporation Composite grafts
EP0257091B1 (en) 1986-02-24 1993-07-28 Robert E. Fischell An intravascular stent and percutaneous insertion system
US4878906A (en) 1986-03-25 1989-11-07 Servetus Partnership Endoprosthesis for repairing a damaged vessel
EP0241838B1 (en) 1986-04-07 1992-04-15 Agency Of Industrial Science And Technology Antithrombogenic material
US4744365A (en) 1986-07-17 1988-05-17 United States Surgical Corporation Two-phase compositions for absorbable surgical devices
US5527337A (en) 1987-06-25 1996-06-18 Duke University Bioabsorbable stent and method of making the same
US4886062A (en) 1987-10-19 1989-12-12 Medtronic, Inc. Intravascular radially expandable stent and method of implant
US4877030A (en) 1988-02-02 1989-10-31 Andreas Beck Device for the widening of blood vessels
US5328471A (en) 1990-02-26 1994-07-12 Endoluminal Therapeutics, Inc. Method and apparatus for treatment of focal disease in hollow tubular organs and other tissue lumens
US5019090A (en) 1988-09-01 1991-05-28 Corvita Corporation Radially expandable endoprosthesis and the like
CA1322628C (en) 1988-10-04 1993-10-05 Richard A. Schatz Expandable intraluminal graft
US4977901A (en) 1988-11-23 1990-12-18 Minnesota Mining And Manufacturing Company Article having non-crosslinked crystallized polymer coatings
CH678393A5 (en) 1989-01-26 1991-09-13 Ulrich Prof Dr Med Sigwart
DE69030811T2 (en) 1989-01-27 1997-10-02 Au Membrane & Biotech Res Inst RECEPTOR MEMBRANES AND SELECTIVE CONTROL OF THE ION FLOW BY IONOPHORES
US5163958A (en) 1989-02-02 1992-11-17 Cordis Corporation Carbon coated tubular endoprosthesis
NZ228382A (en) 1989-03-17 1992-08-26 Carter Holt Harvey Plastic Pro Drug administering coil-like device for insertion in body cavity of animal
US4990158A (en) 1989-05-10 1991-02-05 United States Surgical Corporation Synthetic semiabsorbable tubular prosthesis
US5971954A (en) 1990-01-10 1999-10-26 Rochester Medical Corporation Method of making catheter
ATE120377T1 (en) 1990-02-08 1995-04-15 Howmedica INFLATABLE DILATATOR.
US5545208A (en) 1990-02-28 1996-08-13 Medtronic, Inc. Intralumenal drug eluting prosthesis
US5156623A (en) 1990-04-16 1992-10-20 Olympus Optical Co., Ltd. Sustained release material and method of manufacturing the same
US5123917A (en) 1990-04-27 1992-06-23 Lee Peter Y Expandable intraluminal vascular graft
US6060451A (en) 1990-06-15 2000-05-09 The National Research Council Of Canada Thrombin inhibitors based on the amino acid sequence of hirudin
US5236447A (en) 1990-06-29 1993-08-17 Nissho Corporation Artificial tubular organ
US5342395A (en) 1990-07-06 1994-08-30 American Cyanamid Co. Absorbable surgical repair devices
US5112457A (en) 1990-07-23 1992-05-12 Case Western Reserve University Process for producing hydroxylated plasma-polymerized films and the use of the films for enhancing the compatiblity of biomedical implants
US5455040A (en) 1990-07-26 1995-10-03 Case Western Reserve University Anticoagulant plasma polymer-modified substrate
IL99296A (en) 1990-08-28 1995-12-08 Meadox Medicals Inc Self-supporting woven vascular graft and its preparation
US5163952A (en) 1990-09-14 1992-11-17 Michael Froix Expandable polymeric stent with memory and delivery apparatus and method
DE69116130T2 (en) 1990-10-18 1996-05-15 Ho Young Song SELF-EXPANDING, ENDOVASCULAR DILATATOR
JP2514746B2 (en) * 1990-11-09 1996-07-10 積水化成品工業株式会社 Method for producing porous spherical apatite particles
US5163951A (en) 1990-12-27 1992-11-17 Corvita Corporation Mesh composite graft
CS277367B6 (en) 1990-12-29 1993-01-13 Krajicek Milan Three-layered vascular prosthesis
WO1992015342A1 (en) 1991-03-08 1992-09-17 Keiji Igaki Stent for vessel, structure of holding said stent, and device for mounting said stent
US5356433A (en) 1991-08-13 1994-10-18 Cordis Corporation Biocompatible metal surfaces
US5500013A (en) 1991-10-04 1996-03-19 Scimed Life Systems, Inc. Biodegradable drug delivery vascular stent
US5234457A (en) 1991-10-09 1993-08-10 Boston Scientific Corporation Impregnated stent
US5545408A (en) 1991-10-21 1996-08-13 Peptide Technology Limited Biocompatible implant for the timing of ovulation in mares
US5167614A (en) 1991-10-29 1992-12-01 Medical Engineering Corporation Prostatic stent
US5756476A (en) 1992-01-14 1998-05-26 The United States Of America As Represented By The Department Of Health And Human Services Inhibition of cell proliferation using antisense oligonucleotides
US6537574B1 (en) * 1992-02-11 2003-03-25 Bioform, Inc. Soft tissue augmentation material
DE69332950T2 (en) 1992-03-31 2004-05-13 Boston Scientific Corp., Natick BLOOD VESSEL FILTER
DE4222380A1 (en) 1992-07-08 1994-01-13 Ernst Peter Prof Dr M Strecker Endoprosthesis implantable percutaneously in a patient's body
US5514379A (en) 1992-08-07 1996-05-07 The General Hospital Corporation Hydrogel compositions and methods of use
US5853408A (en) 1992-08-20 1998-12-29 Advanced Cardiovascular Systems, Inc. In-vivo modification of the mechanical properties of surgical devices
US5342621A (en) 1992-09-15 1994-08-30 Advanced Cardiovascular Systems, Inc. Antithrombogenic surface
US5830461A (en) 1992-11-25 1998-11-03 University Of Pittsburgh Of The Commonwealth System Of Higher Education Methods for promoting wound healing and treating transplant-associated vasculopathy
US5342348A (en) 1992-12-04 1994-08-30 Kaplan Aaron V Method and device for treating and enlarging body lumens
EP0604022A1 (en) 1992-12-22 1994-06-29 Advanced Cardiovascular Systems, Inc. Multilayered biodegradable stent and method for its manufacture
US5443458A (en) 1992-12-22 1995-08-22 Advanced Cardiovascular Systems, Inc. Multilayered biodegradable stent and method of manufacture
US5981568A (en) 1993-01-28 1999-11-09 Neorx Corporation Therapeutic inhibitor of vascular smooth muscle cells
JP3583801B2 (en) 1993-03-03 2004-11-04 ボストン サイエンティフィック リミテッド Luminal stents and implants
US5464650A (en) 1993-04-26 1995-11-07 Medtronic, Inc. Intravascular stent and method
EG20321A (en) 1993-07-21 1998-10-31 Otsuka Pharma Co Ltd Medical material and process for producing the same
DE69330132T2 (en) 1993-07-23 2001-11-15 Cook Inc FLEXIBLE STENT WITH A CONFIGURATION MOLDED FROM A MATERIAL SHEET
DK0716610T3 (en) 1993-08-26 2006-09-04 Genetics Inst Llc Human bone morphogenetic proteins for use in neural regeneration
DK0659389T3 (en) 1993-10-20 1999-02-15 Schneider Europ Ag endoprosthesis
US5723004A (en) 1993-10-21 1998-03-03 Corvita Corporation Expandable supportive endoluminal grafts
SE501288C2 (en) 1993-11-30 1995-01-09 Corimed Gmbh Process for preparing ceramic implant material, preferably hydroxylapatite having ceramic implant material
US5626611A (en) 1994-02-10 1997-05-06 United States Surgical Corporation Composite bioabsorbable materials and surgical articles made therefrom
AU704549B2 (en) 1994-03-18 1999-04-29 Lynx Therapeutics, Inc. Oligonucleotide N3'-P5' phosphoramidates: synthesis and compounds; hybridization and nuclease resistance properties
US5626861A (en) 1994-04-01 1997-05-06 Massachusetts Institute Of Technology Polymeric-hydroxyapatite bone composite
US5693085A (en) 1994-04-29 1997-12-02 Scimed Life Systems, Inc. Stent with collagen
US5629077A (en) 1994-06-27 1997-05-13 Advanced Cardiovascular Systems, Inc. Biodegradable mesh and film stent
US5670558A (en) 1994-07-07 1997-09-23 Terumo Kabushiki Kaisha Medical instruments that exhibit surface lubricity when wetted
US6007845A (en) 1994-07-22 1999-12-28 Massachusetts Institute Of Technology Nanoparticles and microparticles of non-linear hydrophilic-hydrophobic multiblock copolymers
US5554120A (en) 1994-07-25 1996-09-10 Advanced Cardiovascular Systems, Inc. Polymer blends for use in making medical devices including catheters and balloons for dilatation catheters
US6015429A (en) 1994-09-08 2000-01-18 Gore Enterprise Holdings, Inc. Procedures for introducing stents and stent-grafts
US5578073A (en) 1994-09-16 1996-11-26 Ramot Of Tel Aviv University Thromboresistant surface treatment for biomaterials
US5649977A (en) 1994-09-22 1997-07-22 Advanced Cardiovascular Systems, Inc. Metal reinforced polymer stent
US5765682A (en) 1994-10-13 1998-06-16 Menlo Care, Inc. Restrictive package for expandable or shape memory medical devices and method of preventing premature change of same
US5836964A (en) 1996-10-30 1998-11-17 Medinol Ltd. Stent fabrication method
IL115755A0 (en) 1994-10-27 1996-01-19 Medinol Ltd X-ray visible stent
CA2301351C (en) 1994-11-28 2002-01-22 Advanced Cardiovascular Systems, Inc. Method and apparatus for direct laser cutting of metal stents
US5637113A (en) 1994-12-13 1997-06-10 Advanced Cardiovascular Systems, Inc. Polymer film for wrapping a stent structure
US5919570A (en) 1995-02-01 1999-07-06 Schneider Inc. Slippery, tenaciously adhering hydrogel coatings containing a polyurethane-urea polymer hydrogel commingled with a poly(N-vinylpyrrolidone) polymer hydrogel, coated polymer and metal substrate materials, and coated medical devices
US6017577A (en) 1995-02-01 2000-01-25 Schneider (Usa) Inc. Slippery, tenaciously adhering hydrophilic polyurethane hydrogel coatings, coated polymer substrate materials, and coated medical devices
US6120536A (en) 1995-04-19 2000-09-19 Schneider (Usa) Inc. Medical devices with long term non-thrombogenic coatings
US5837313A (en) 1995-04-19 1998-11-17 Schneider (Usa) Inc Drug release stent coating process
US5954744A (en) 1995-06-06 1999-09-21 Quanam Medical Corporation Intravascular stent
US5609629A (en) 1995-06-07 1997-03-11 Med Institute, Inc. Coated implantable medical device
US6129761A (en) 1995-06-07 2000-10-10 Reprogenesis, Inc. Injectable hydrogel compositions
US5667767A (en) 1995-07-27 1997-09-16 Micro Therapeutics, Inc. Compositions for use in embolizing blood vessels
GB9611437D0 (en) 1995-08-03 1996-08-07 Secr Defence Biomaterial
US5830879A (en) 1995-10-02 1998-11-03 St. Elizabeth's Medical Center Of Boston, Inc. Treatment of vascular injury using vascular endothelial growth factor
US6048964A (en) 1995-12-12 2000-04-11 Stryker Corporation Compositions and therapeutic methods using morphogenic proteins and stimulatory factors
DK2111876T3 (en) 1995-12-18 2011-12-12 Angiodevice Internat Gmbh Crosslinked polymer preparations and methods for their use
ATE290832T1 (en) 1996-01-05 2005-04-15 Medtronic Inc EXPANDABLE ENDOLUMINAL PROSTHESES
US6150630A (en) 1996-01-11 2000-11-21 The Regents Of The University Of California Laser machining of explosives
EP1011889B1 (en) 1996-01-30 2002-10-30 Medtronic, Inc. Articles for and methods of making stents
JP2000509014A (en) 1996-03-11 2000-07-18 フォーカル,インコーポレイテッド Polymer delivery of radionuclides and radiopharmaceuticals
US6241760B1 (en) 1996-04-26 2001-06-05 G. David Jang Intravascular stent
US6071266A (en) 1996-04-26 2000-06-06 Kelley; Donald W. Lubricious medical devices
US6592617B2 (en) 1996-04-30 2003-07-15 Boston Scientific Scimed, Inc. Three-dimensional braided covered stent
US5670161A (en) * 1996-05-28 1997-09-23 Healy; Kevin E. Biodegradable stent
US5830178A (en) 1996-10-11 1998-11-03 Micro Therapeutics, Inc. Methods for embolizing vascular sites with an emboilizing composition comprising dimethylsulfoxide
US5800516A (en) 1996-08-08 1998-09-01 Cordis Corporation Deployable and retrievable shape memory stent/tube and method
US6344271B1 (en) 1998-11-06 2002-02-05 Nanoenergy Corporation Materials and products using nanostructured non-stoichiometric substances
JP3773335B2 (en) 1996-10-04 2006-05-10 ダイセル化学工業株式会社 Biodegradable aliphatic polyester resin-starch composition
US6387121B1 (en) 1996-10-21 2002-05-14 Inflow Dynamics Inc. Vascular and endoluminal stents with improved coatings
US5833651A (en) 1996-11-08 1998-11-10 Medtronic, Inc. Therapeutic intraluminal stents
IT1289728B1 (en) 1996-12-10 1998-10-16 Sorin Biomedica Cardio Spa SYSTEM AND EQUIPMENT DEVICE THAT INCLUDES IT
US5980972A (en) 1996-12-20 1999-11-09 Schneider (Usa) Inc Method of applying drug-release coatings
US5906759A (en) 1996-12-26 1999-05-25 Medinol Ltd. Stent forming apparatus with stent deforming blades
IT1291001B1 (en) 1997-01-09 1998-12-14 Sorin Biomedica Cardio Spa ANGIOPLASTIC STENT AND ITS PRODUCTION PROCESS
US6159951A (en) 1997-02-13 2000-12-12 Ribozyme Pharmaceuticals Inc. 2'-O-amino-containing nucleoside analogs and polynucleotides
US6582472B2 (en) 1997-02-26 2003-06-24 Applied Medical Resources Corporation Kinetic stent
US6210715B1 (en) 1997-04-01 2001-04-03 Cap Biotechnology, Inc. Calcium phosphate microcarriers and microspheres
US8172897B2 (en) 1997-04-15 2012-05-08 Advanced Cardiovascular Systems, Inc. Polymer and metal composite implantable medical devices
US6240616B1 (en) 1997-04-15 2001-06-05 Advanced Cardiovascular Systems, Inc. Method of manufacturing a medicated porous metal prosthesis
US6273913B1 (en) 1997-04-18 2001-08-14 Cordis Corporation Modified stent useful for delivery of drugs along stent strut
FI103715B (en) 1997-04-21 1999-08-31 Vivoxid Oy New composite and its use
US5741327A (en) 1997-05-06 1998-04-21 Global Therapeutics, Inc. Surgical stent featuring radiopaque markers
US6303901B1 (en) 1997-05-20 2001-10-16 The Regents Of The University Of California Method to reduce damage to backing plate
US6056993A (en) 1997-05-30 2000-05-02 Schneider (Usa) Inc. Porous protheses and methods for making the same wherein the protheses are formed by spraying water soluble and water insoluble fibers onto a rotating mandrel
US5980928A (en) 1997-07-29 1999-11-09 Terry; Paul B. Implant for preventing conjunctivitis in cattle
US5980564A (en) 1997-08-01 1999-11-09 Schneider (Usa) Inc. Bioabsorbable implantable endoprosthesis with reservoir
US6245103B1 (en) 1997-08-01 2001-06-12 Schneider (Usa) Inc Bioabsorbable self-expanding stent
US6340367B1 (en) 1997-08-01 2002-01-22 Boston Scientific Scimed, Inc. Radiopaque markers and methods of using the same
US6121027A (en) 1997-08-15 2000-09-19 Surmodics, Inc. Polybifunctional reagent having a polymeric backbone and photoreactive moieties and bioactive groups
US6117979A (en) 1997-08-18 2000-09-12 Medtronic, Inc. Process for making a bioprosthetic device and implants produced therefrom
US6129928A (en) 1997-09-05 2000-10-10 Icet, Inc. Biomimetic calcium phosphate implant coatings and methods for making the same
US6284333B1 (en) 1997-09-10 2001-09-04 Scimed Life Systems, Inc. Medical devices made from polymer blends containing low melting temperature liquid crystal polymers
WO1999016871A2 (en) 1997-09-22 1999-04-08 Max-Planck-Gesellschaft Zur Forderung Der Wissensc Nucleic acid catalysts with endonuclease activity
DE69838256T2 (en) 1997-09-24 2008-05-15 Med Institute, Inc., West Lafayette RADIAL EXPANDABLE STENT
US5976182A (en) 1997-10-03 1999-11-02 Advanced Cardiovascular Systems, Inc. Balloon-expandable, crush-resistant locking stent and method of loading the same
DE19881727D2 (en) 1997-11-24 2001-01-04 Herbert P Jennissen Process for immobilizing mediator molecules on inorganic and metallic implant materials
US6093463A (en) 1997-12-12 2000-07-25 Intella Interventional Systems, Inc. Medical devices made from improved polymer blends
US5957975A (en) 1997-12-15 1999-09-28 The Cleveland Clinic Foundation Stent having a programmed pattern of in vivo degradation
US6626939B1 (en) 1997-12-18 2003-09-30 Boston Scientific Scimed, Inc. Stent-graft with bioabsorbable structural support
US5986169A (en) 1997-12-31 1999-11-16 Biorthex Inc. Porous nickel-titanium alloy article
US6224626B1 (en) 1998-02-17 2001-05-01 Md3, Inc. Ultra-thin expandable stent
DK1062278T3 (en) 1998-02-23 2006-09-25 Mnemoscience Gmbh Polymers with shape memory
RU2215542C2 (en) 1998-02-23 2003-11-10 Массачусетс Инститьют Оф Текнолоджи Biodecomposing polymers able recovery of form
US5938697A (en) 1998-03-04 1999-08-17 Scimed Life Systems, Inc. Stent having variable properties
US6110188A (en) 1998-03-09 2000-08-29 Corvascular, Inc. Anastomosis method
US6113629A (en) 1998-05-01 2000-09-05 Micrus Corporation Hydrogel for the therapeutic treatment of aneurysms
US6083258A (en) 1998-05-28 2000-07-04 Yadav; Jay S. Locking stent
DE19856983A1 (en) 1998-06-25 1999-12-30 Biotronik Mess & Therapieg Implantable, bioresorbable vascular wall support, in particular coronary stent
EP0966979B1 (en) 1998-06-25 2006-03-08 Biotronik AG Implantable bioresorbable support for the vascular walls, in particular coronary stent
US6153252A (en) 1998-06-30 2000-11-28 Ethicon, Inc. Process for coating stents
WO2000023123A1 (en) 1998-10-19 2000-04-27 Synthes Ag Chur Hardenable ceramic hydraulic cement
DE69822470T2 (en) 1998-11-12 2005-01-20 Takiron Co. Ltd. Biodegradable absorbable shape memory material
US6165486A (en) 1998-11-19 2000-12-26 Carnegie Mellon University Biocompatible compositions and methods of using same
US6125523A (en) 1998-11-20 2000-10-03 Advanced Cardiovascular Systems, Inc. Stent crimping tool and method of use
US6350277B1 (en) 1999-01-15 2002-02-26 Scimed Life Systems, Inc. Stents with temporary retaining bands
US6395029B1 (en) 1999-01-19 2002-05-28 The Children's Hospital Of Philadelphia Sustained delivery of polyionic bioactive agents
DE60017363T2 (en) 1999-02-02 2006-03-02 Wright Medical Technology Inc., Arlington CONTROLLED RELEASE OF A COMPOSITE MATERIAL
US6066156A (en) 1999-03-11 2000-05-23 Advanced Cardiovascular Systems, Inc. Temperature activated adhesive for releasably attaching stents to balloons
US6667049B2 (en) 1999-06-14 2003-12-23 Ethicon, Inc. Relic process for producing bioresorbable ceramic tissue scaffolds
US6312459B1 (en) 1999-06-30 2001-11-06 Advanced Cardiovascular Systems, Inc. Stent design for use in small vessels
AUPQ170799A0 (en) 1999-07-20 1999-08-12 Cardiac Crc Nominees Pty Limited Shape memory polyurethane or polyurethane-urea polymers
US6569193B1 (en) 1999-07-22 2003-05-27 Advanced Cardiovascular Systems, Inc. Tapered self-expanding stent
US6479565B1 (en) 1999-08-16 2002-11-12 Harold R. Stanley Bioactive ceramic cement
US6379381B1 (en) 1999-09-03 2002-04-30 Advanced Cardiovascular Systems, Inc. Porous prosthesis and a method of depositing substances into the pores
JP4172883B2 (en) 1999-09-08 2008-10-29 Hoya株式会社 Drug sustained release carrier and method for producing drug sustained release carrier
WO2001026584A1 (en) 1999-10-14 2001-04-19 United Stenting, Inc. Stents with multilayered struts
US7226475B2 (en) 1999-11-09 2007-06-05 Boston Scientific Scimed, Inc. Stent with variable properties
WO2001035928A1 (en) 1999-11-17 2001-05-25 Microchips, Inc. Microfabricated devices for the delivery of molecules into a carrier fluid
US7947069B2 (en) 1999-11-24 2011-05-24 University Of Washington Medical devices comprising small fiber biomaterials, and methods of use
US20040009228A1 (en) 1999-11-30 2004-01-15 Pertti Tormala Bioabsorbable drug delivery system for local treatment and prevention of infections
US6554854B1 (en) 1999-12-10 2003-04-29 Scimed Life Systems, Inc. Process for laser joining dissimilar metals and endoluminal stent with radiopaque marker produced thereby
US6494908B1 (en) 1999-12-22 2002-12-17 Ethicon, Inc. Removable stent for body lumens
US6338739B1 (en) 1999-12-22 2002-01-15 Ethicon, Inc. Biodegradable stent
US6375826B1 (en) 2000-02-14 2002-04-23 Advanced Cardiovascular Systems, Inc. Electro-polishing fixture and electrolyte solution for polishing stents and method
DE60112950T2 (en) 2000-03-18 2006-05-18 Polyzenix Gmbh Dental implants with antibacterial properties
EP1153621A1 (en) 2000-05-12 2001-11-14 MERCK PATENT GmbH Biocements based on a mixture of TCP-PHA with improved compressive strength
US6395326B1 (en) 2000-05-31 2002-05-28 Advanced Cardiovascular Systems, Inc. Apparatus and method for depositing a coating onto a surface of a prosthesis
IL137090A (en) 2000-06-29 2010-04-15 Pentech Medical Devices Ltd Polymeric stent
US6569191B1 (en) 2000-07-27 2003-05-27 Bionx Implants, Inc. Self-expanding stent with enhanced radial expansion and shape memory
US6574851B1 (en) 2000-07-31 2003-06-10 Advanced Cardiovascular Systems, Inc. Stent made by rotational molding or centrifugal casting and method for making the same
US6485512B1 (en) 2000-09-27 2002-11-26 Advanced Cardiovascular Systems, Inc. Two-stage light curable stent and delivery system
US6746773B2 (en) 2000-09-29 2004-06-08 Ethicon, Inc. Coatings for medical devices
US20020111590A1 (en) 2000-09-29 2002-08-15 Davila Luis A. Medical devices, drug coatings and methods for maintaining the drug coatings thereon
US6492615B1 (en) 2000-10-12 2002-12-10 Scimed Life Systems, Inc. Laser polishing of medical devices
US6664335B2 (en) 2000-11-30 2003-12-16 Cardiac Pacemakers, Inc. Polyurethane elastomer article with “shape memory” and medical devices therefrom
US20020103526A1 (en) 2000-12-15 2002-08-01 Tom Steinke Protective coating for stent
US6565599B1 (en) 2000-12-28 2003-05-20 Advanced Cardiovascular Systems, Inc. Hybrid stent
US6563080B2 (en) 2001-02-15 2003-05-13 Scimed Life Systems, Inc. Laser cutting of stents and other medical devices
US6540777B2 (en) 2001-02-15 2003-04-01 Scimed Life Systems, Inc. Locking stent
US6764505B1 (en) 2001-04-12 2004-07-20 Advanced Cardiovascular Systems, Inc. Variable surface area stent
US6730324B2 (en) 2001-04-20 2004-05-04 The University Of British Columbia Biofunctional hydroxyapatite coatings and microspheres for in-situ drug encapsulation
US7241856B2 (en) 2003-06-02 2007-07-10 Pentron Clinical Technologies Llc Dental resins, dental composite materials, and method of manufacture thereof
US6585755B2 (en) 2001-06-29 2003-07-01 Advanced Cardiovascular Polymeric stent suitable for imaging by MRI and fluoroscopy
JP2005503865A (en) 2001-09-28 2005-02-10 ボストン サイエンティフィック リミテッド Medical device comprising nanomaterial and treatment method using the same
US6620202B2 (en) * 2001-10-16 2003-09-16 Scimed Life Systems, Inc. Medical stent with variable coil and related methods
CN1301757C (en) * 2001-11-27 2007-02-28 多喜兰株式会社 Implant material and process for producing the same
US20030105530A1 (en) 2001-12-04 2003-06-05 Inion Ltd. Biodegradable implant and method for manufacturing one
US6752826B2 (en) 2001-12-14 2004-06-22 Thoratec Corporation Layered stent-graft and methods of making the same
US20030187495A1 (en) 2002-04-01 2003-10-02 Cully Edward H. Endoluminal devices, embolic filters, methods of manufacture and use
US7270675B2 (en) 2002-05-10 2007-09-18 Cordis Corporation Method of forming a tubular membrane on a structural frame
US20030236563A1 (en) 2002-06-20 2003-12-25 Dan Fifer Stent delivery catheter with retention bands
US7141063B2 (en) 2002-08-06 2006-11-28 Icon Medical Corp. Stent with micro-latching hinge joints
US6818063B1 (en) 2002-09-24 2004-11-16 Advanced Cardiovascular Systems, Inc. Stent mandrel fixture and method for minimizing coating defects
US6966990B2 (en) 2002-10-11 2005-11-22 Ferro Corporation Composite particles and method for preparing
US6992127B2 (en) * 2002-11-25 2006-01-31 Ast Products, Inc. Polymeric coatings containing a pH buffer agent
US7776926B1 (en) 2002-12-11 2010-08-17 Advanced Cardiovascular Systems, Inc. Biocompatible coating for implantable medical devices
JP2004189868A (en) 2002-12-11 2004-07-08 Maruo Calcium Co Ltd Thermoplastic resin composition and molded product
US7455687B2 (en) 2002-12-30 2008-11-25 Advanced Cardiovascular Systems, Inc. Polymer link hybrid stent
US20040143317A1 (en) 2003-01-17 2004-07-22 Stinson Jonathan S. Medical devices
US20040167610A1 (en) 2003-02-26 2004-08-26 Fleming James A. Locking stent
US6894145B2 (en) 2003-06-16 2005-05-17 Organic Vision Inc. Methods to purify polymers
WO2005051229A2 (en) 2003-11-24 2005-06-09 Avantec Vascular Corporation Devices delivering therapeutic agents and methods regarding the same
JP4587198B2 (en) 2003-12-05 2010-11-24 丸尾カルシウム株式会社 Foamed resin molding
US20050181015A1 (en) 2004-02-12 2005-08-18 Sheng-Ping (Samuel) Zhong Layered silicate nanoparticles for controlled delivery of therapeutic agents from medical articles
US7390335B2 (en) * 2004-04-06 2008-06-24 American Dental Association Foundation Nanostructured bioactive materials prepared by spray drying techniques
US7316738B2 (en) 2004-10-08 2008-01-08 Phibro-Tech, Inc. Milled submicron chlorothalonil with narrow particle size distribution, and uses thereof
US7785615B2 (en) 2004-05-28 2010-08-31 Cordis Corporation Biodegradable medical implant with encapsulated buffering agent
US20050278929A1 (en) 2004-06-16 2005-12-22 National Taipei University Technology Process of manufacturing stent with therapeutic function in the human body
US20070224286A1 (en) * 2004-06-25 2007-09-27 Kutty Thundyil Raman N Process for Preparing Calcium Phosphate Self-Setting Bone Cement, the Cement So Prepared and Uses Thereof
EP1808411B1 (en) * 2004-09-14 2013-10-23 Japan Science and Technology Agency Ceramic particle group and method for production thereof and use thereof
US7862835B2 (en) * 2004-10-27 2011-01-04 Boston Scientific Scimed, Inc. Method of manufacturing a medical device having a porous coating thereon
US8535702B2 (en) 2005-02-01 2013-09-17 Boston Scientific Scimed, Inc. Medical devices having porous polymeric regions for controlled drug delivery and regulated biocompatibility
JP5247984B2 (en) 2005-02-10 2013-07-24 コーディス・コーポレイション Biodegradable medical device with enhanced mechanical strength and pharmacological function
US20060199876A1 (en) 2005-03-04 2006-09-07 The University Of British Columbia Bioceramic composite coatings and process for making same
US20060270752A1 (en) * 2005-05-26 2006-11-30 Ada Foundation Dental releasing materials
US9248034B2 (en) * 2005-08-23 2016-02-02 Advanced Cardiovascular Systems, Inc. Controlled disintegrating implantable medical devices
US8585753B2 (en) 2006-03-04 2013-11-19 John James Scanlon Fibrillated biodegradable prosthesis
US20090113641A1 (en) * 2006-03-06 2009-05-07 Yoshiki Akatani Ink for Inkjet Textile Printing and an Inkjet Textile Printing Method Using the Same
US7964210B2 (en) 2006-03-31 2011-06-21 Abbott Cardiovascular Systems Inc. Degradable polymeric implantable medical devices with a continuous phase and discrete phase
US7959940B2 (en) * 2006-05-30 2011-06-14 Advanced Cardiovascular Systems, Inc. Polymer-bioceramic composite implantable medical devices
US9089627B2 (en) 2006-07-11 2015-07-28 Abbott Cardiovascular Systems Inc. Stent fabricated from polymer composite toughened by a dispersed phase
US8202528B2 (en) 2007-06-05 2012-06-19 Abbott Cardiovascular Systems Inc. Implantable medical devices with elastomeric block copolymer coatings

Patent Citations (101)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4321711A (en) * 1978-10-18 1982-03-30 Sumitomo Electric Industries, Ltd. Vascular prosthesis
US4902289A (en) * 1982-04-19 1990-02-20 Massachusetts Institute Of Technology Multilayer bioreplaceable blood vessel prosthesis
US4656083A (en) * 1983-08-01 1987-04-07 Washington Research Foundation Plasma gas discharge treatment for improving the biocompatibility of biomaterials
US5197977A (en) * 1984-01-30 1993-03-30 Meadox Medicals, Inc. Drug delivery collagen-impregnated synthetic vascular graft
US4633873A (en) * 1984-04-26 1987-01-06 American Cyanamid Company Surgical repair mesh
US4732152A (en) * 1984-12-05 1988-03-22 Medinvent S.A. Device for implantation and a method of implantation in a vessel using such device
US4718907A (en) * 1985-06-20 1988-01-12 Atrium Medical Corporation Vascular prosthesis having fluorinated coating with varying F/C ratio
US4818559A (en) * 1985-08-08 1989-04-04 Sumitomo Chemical Company, Limited Method for producing endosseous implants
US4733665C2 (en) * 1985-11-07 2002-01-29 Expandable Grafts Partnership Expandable intraluminal graft and method and apparatus for implanting an expandable intraluminal graft
US4733665A (en) * 1985-11-07 1988-03-29 Expandable Grafts Partnership Expandable intraluminal graft, and method and apparatus for implanting an expandable intraluminal graft
US4739762B1 (en) * 1985-11-07 1998-10-27 Expandable Grafts Partnership Expandable intraluminal graft and method and apparatus for implanting an expandable intraluminal graft
US4733665B1 (en) * 1985-11-07 1994-01-11 Expandable Grafts Partnership Expandable intraluminal graft,and method and apparatus for implanting an expandable intraluminal graft
US4739762A (en) * 1985-11-07 1988-04-26 Expandable Grafts Partnership Expandable intraluminal graft, and method and apparatus for implanting an expandable intraluminal graft
US4740207A (en) * 1986-09-10 1988-04-26 Kreamer Jeffry W Intralumenal graft
US4723549A (en) * 1986-09-18 1988-02-09 Wholey Mark H Method and apparatus for dilating blood vessels
US4722335A (en) * 1986-10-20 1988-02-02 Vilasi Joseph A Expandable endotracheal tube
US4800882A (en) * 1987-03-13 1989-01-31 Cook Incorporated Endovascular stent and delivery system
US4816339A (en) * 1987-04-28 1989-03-28 Baxter International Inc. Multi-layered poly(tetrafluoroethylene)/elastomer materials useful for in vivo implantation
US5306286A (en) * 1987-06-25 1994-04-26 Duke University Absorbable stent
US5192311A (en) * 1988-04-25 1993-03-09 Angeion Corporation Medical implant and method of making
US4994298A (en) * 1988-06-07 1991-02-19 Biogold Inc. Method of making a biocompatible prosthesis
US5502158A (en) * 1988-08-08 1996-03-26 Ecopol, Llc Degradable polymer composition
US5085629A (en) * 1988-10-06 1992-02-04 Medical Engineering Corporation Biodegradable stent
US5289831A (en) * 1989-03-09 1994-03-01 Vance Products Incorporated Surface-treated stent, catheter, cannula, and the like
US5108755A (en) * 1989-04-27 1992-04-28 Sri International Biodegradable composites for internal medical use
US5100429A (en) * 1989-04-28 1992-03-31 C. R. Bard, Inc. Endovascular stent and delivery system
US5407609A (en) * 1989-05-04 1995-04-18 Southern Research Institute Microencapsulation process and products therefrom
US5084065A (en) * 1989-07-10 1992-01-28 Corvita Corporation Reinforced graft assembly
US5290271A (en) * 1990-05-14 1994-03-01 Jernberg Gary R Surgical implant and method for controlled release of chemotherapeutic agents
US5279594A (en) * 1990-05-23 1994-01-18 Jackson Richard R Intubation devices with local anesthetic effect for medical use
US5385580A (en) * 1990-08-28 1995-01-31 Meadox Medicals, Inc. Self-supporting woven vascular graft
US5607467A (en) * 1990-09-14 1997-03-04 Froix; Michael Expandable polymeric stent with memory and delivery apparatus and method
US5108417A (en) * 1990-09-14 1992-04-28 Interface Biomedical Laboratories Corp. Anti-turbulent, anti-thrombogenic intravascular stent
US5104410A (en) * 1990-10-22 1992-04-14 Intermedics Orthopedics, Inc Surgical implant having multiple layers of sintered porous coating and method
US5711763A (en) * 1991-02-20 1998-01-27 Tdk Corporation Composite biological implant of a ceramic material in a metal substrate
US5282860A (en) * 1991-10-16 1994-02-01 Olympus Optical Co., Ltd. Stent tube for medical use
US5593434A (en) * 1992-01-31 1997-01-14 Advanced Cardiovascular Systems, Inc. Stent capable of attachment within a body lumen
US5858746A (en) * 1992-04-20 1999-01-12 Board Of Regents, The University Of Texas System Gels for encapsulation of biological materials
US5306294A (en) * 1992-08-05 1994-04-26 Ultrasonic Sensing And Monitoring Systems, Inc. Stent construction of rolled configuration
US5383925A (en) * 1992-09-14 1995-01-24 Meadox Medicals, Inc. Three-dimensional braided soft tissue prosthesis
US5733925A (en) * 1993-01-28 1998-03-31 Neorx Corporation Therapeutic inhibitor of vascular smooth muscle cells
US5733564A (en) * 1993-04-14 1998-03-31 Leiras Oy Method of treating endo-osteal materials with a bisphosphonate solution
US5618299A (en) * 1993-04-23 1997-04-08 Advanced Cardiovascular Systems, Inc. Ratcheting stent
US5716981A (en) * 1993-07-19 1998-02-10 Angiogenesis Technologies, Inc. Anti-angiogenic compositions and methods of use
US5389106A (en) * 1993-10-29 1995-02-14 Numed, Inc. Impermeable expandable intravascular stent
US5599301A (en) * 1993-11-22 1997-02-04 Advanced Cardiovascular Systems, Inc. Motor control system for an automatic catheter inflation system
US5725549A (en) * 1994-03-11 1998-03-10 Advanced Cardiovascular Systems, Inc. Coiled stent with locking ends
US5726297A (en) * 1994-03-18 1998-03-10 Lynx Therapeutics, Inc. Oligodeoxyribonucleotide N3' P5' phosphoramidates
US5591607A (en) * 1994-03-18 1997-01-07 Lynx Therapeutics, Inc. Oligonucleotide N3→P5' phosphoramidates: triplex DNA formation
US6169170B1 (en) * 1994-03-18 2001-01-02 Lynx Therapeutics, Inc. Oligonucleotide N3′→N5′Phosphoramidate Duplexes
US5599922A (en) * 1994-03-18 1997-02-04 Lynx Therapeutics, Inc. Oligonucleotide N3'-P5' phosphoramidates: hybridization and nuclease resistance properties
US5399666A (en) * 1994-04-21 1995-03-21 E. I. Du Pont De Nemours And Company Easily degradable star-block copolymers
US5874109A (en) * 1994-07-27 1999-02-23 The Trustees Of The University Of Pennsylvania Incorporation of biological molecules into bioactive glasses
US5593403A (en) * 1994-09-14 1997-01-14 Scimed Life Systems Inc. Method for modifying a stent in an implanted site
US5879713A (en) * 1994-10-12 1999-03-09 Focal, Inc. Targeted delivery via biodegradable polymers
US5707385A (en) * 1994-11-16 1998-01-13 Advanced Cardiovascular Systems, Inc. Drug loaded elastic membrane and method for delivery
US6171609B1 (en) * 1995-02-15 2001-01-09 Neorx Corporation Therapeutic inhibitor of vascular smooth muscle cells
US5876743A (en) * 1995-03-21 1999-03-02 Den-Mat Corporation Biocompatible adhesion in tissue repair
US5605696A (en) * 1995-03-30 1997-02-25 Advanced Cardiovascular Systems, Inc. Drug loaded polymeric material and method of manufacture
US20020004101A1 (en) * 1995-04-19 2002-01-10 Schneider (Usa) Inc. Drug coating with topcoat
US5855612A (en) * 1995-05-12 1999-01-05 Ohta Inc. Biocompatible titanium implant
US5591199A (en) * 1995-06-07 1997-01-07 Porter; Christopher H. Curable fiber composite stent and delivery system
US5865814A (en) * 1995-06-07 1999-02-02 Medtronic, Inc. Blood contacting medical device and method
US5873904A (en) * 1995-06-07 1999-02-23 Cook Incorporated Silver implantable medical device
US5888533A (en) * 1995-10-27 1999-03-30 Atrix Laboratories, Inc. Non-polymeric sustained release delivery system
US5607442A (en) * 1995-11-13 1997-03-04 Isostent, Inc. Stent with improved radiopacity and appearance characteristics
US5733326A (en) * 1996-05-28 1998-03-31 Cordis Corporation Composite material endoprosthesis
US5897955A (en) * 1996-06-03 1999-04-27 Gore Hybrid Technologies, Inc. Materials and methods for the immobilization of bioactive species onto polymeric substrates
US5874165A (en) * 1996-06-03 1999-02-23 Gore Enterprise Holdings, Inc. Materials and method for the immobilization of bioactive species onto polymeric subtrates
US5855618A (en) * 1996-09-13 1999-01-05 Meadox Medicals, Inc. Polyurethanes grafted with polyethylene oxide chains containing covalently bonded heparin
US6676697B1 (en) * 1996-09-19 2004-01-13 Medinol Ltd. Stent with variable features to optimize support and method of making such stent
US5868781A (en) * 1996-10-22 1999-02-09 Scimed Life Systems, Inc. Locking stent
US5741881A (en) * 1996-11-25 1998-04-21 Meadox Medicals, Inc. Process for preparing covalently bound-heparin containing polyurethane-peo-heparin coating compositions
US5728751A (en) * 1996-11-25 1998-03-17 Meadox Medicals, Inc. Bonding bio-active materials to substrate surfaces
US5877263A (en) * 1996-11-25 1999-03-02 Meadox Medicals, Inc. Process for preparing polymer coatings grafted with polyethylene oxide chains containing covalently bonded bio-active agents
US5733330A (en) * 1997-01-13 1998-03-31 Advanced Cardiovascular Systems, Inc. Balloon-expandable, crush-resistant locking stent
US5874101A (en) * 1997-04-14 1999-02-23 Usbiomaterials Corp. Bioactive-gel compositions and methods
US6042875A (en) * 1997-04-30 2000-03-28 Schneider (Usa) Inc. Drug-releasing coatings for medical devices
US5891192A (en) * 1997-05-22 1999-04-06 The Regents Of The University Of California Ion-implanted protein-coated intralumenal implants
US20020004060A1 (en) * 1997-07-18 2002-01-10 Bernd Heublein Metallic implant which is degradable in vivo
US6174330B1 (en) * 1997-08-01 2001-01-16 Schneider (Usa) Inc Bioabsorbable marker having radiopaque constituents
US6010445A (en) * 1997-09-11 2000-01-04 Implant Sciences Corporation Radioactive medical device and process
US6015541A (en) * 1997-11-03 2000-01-18 Micro Therapeutics, Inc. Radioactive embolizing compositions
US6511748B1 (en) * 1998-01-06 2003-01-28 Aderans Research Institute, Inc. Bioabsorbable fibers and reinforced composites produced therefrom
US6709379B1 (en) * 1998-11-02 2004-03-23 Alcove Surfaces Gmbh Implant with cavities containing therapeutic agents
US6187045B1 (en) * 1999-02-10 2001-02-13 Thomas K. Fehring Enhanced biocompatible implants and alloys
US6183505B1 (en) * 1999-03-11 2001-02-06 Medtronic Ave, Inc. Method of stent retention to a delivery catheter balloon-braided retainers
US6689823B1 (en) * 1999-03-31 2004-02-10 The Brigham And Women's Hospital, Inc. Nanocomposite surgical materials and method of producing them
US6177523B1 (en) * 1999-07-14 2001-01-23 Cardiotech International, Inc. Functionalized polyurethanes
US6706273B1 (en) * 1999-08-14 2004-03-16 Ivoclar Vivadent Ag Composition for implantation into the human and animal body
US6689375B1 (en) * 1999-11-09 2004-02-10 Coripharm Medizinprodukte Gmbh & Co. Kg Resorbable bone implant material and method for producing the same
US20020002399A1 (en) * 1999-12-22 2002-01-03 Huxel Shawn Thayer Removable stent for body lumens
US6537589B1 (en) * 2000-04-03 2003-03-25 Kyung Won Medical Co., Ltd. Calcium phosphate artificial bone as osteoconductive and biodegradable bone substitute material
US6527801B1 (en) * 2000-04-13 2003-03-04 Advanced Cardiovascular Systems, Inc. Biodegradable drug delivery material for stent
US6517888B1 (en) * 2000-11-28 2003-02-11 Scimed Life Systems, Inc. Method for manufacturing a medical device having a coated portion by laser ablation
US20030033001A1 (en) * 2001-02-27 2003-02-13 Keiji Igaki Stent holding member and stent feeding system
US6679980B1 (en) * 2001-06-13 2004-01-20 Advanced Cardiovascular Systems, Inc. Apparatus for electropolishing a stent
US6695920B1 (en) * 2001-06-27 2004-02-24 Advanced Cardiovascular Systems, Inc. Mandrel for supporting a stent and a method of using the mandrel to coat a stent
US20070026069A1 (en) * 2003-03-28 2007-02-01 Shastri Venkatram P Biommetic hierarchies using functionalized nanoparticles as building blocks
US6846323B2 (en) * 2003-05-15 2005-01-25 Advanced Cardiovascular Systems, Inc. Intravascular stent
US20080063685A1 (en) * 2006-09-13 2008-03-13 Yunbing Wang Degradable polymeric implantable medical devices with continuous phase and discrete phase

Cited By (43)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US9381275B2 (en) 2006-09-25 2016-07-05 Orthovita, Inc. Bioactive load-bearing composites
US10195308B2 (en) 2006-09-25 2019-02-05 Orthovita, Inc. Bioactive load-bearing composites
US9119905B2 (en) * 2007-01-19 2015-09-01 Elixir Medical Corporation Biodegradable endoprostheses and methods for their fabrication
US20150025619A1 (en) * 2007-01-19 2015-01-22 Elixir Medical Corporation Biodegradable endoprostheses and methods for their fabrication
US9566371B2 (en) 2007-01-19 2017-02-14 Elixir Medical Corporation Biodegradable endoprostheses and methods for their fabrication
US9144487B2 (en) 2007-06-11 2015-09-29 Abbott Cardiovascular Systems Inc. Polymer-bioceramic composite medical devices with bioceramic particles having grafted polymers
US8425591B1 (en) 2007-06-11 2013-04-23 Abbott Cardiovascular Systems Inc. Methods of forming polymer-bioceramic composite medical devices with bioceramic particles
US20090088834A1 (en) * 2007-09-28 2009-04-02 Abbott Cardiovascular Systems Inc. Stent formed from bioerodible metal-bioceramic composite
US8998978B2 (en) 2007-09-28 2015-04-07 Abbott Cardiovascular Systems Inc. Stent formed from bioerodible metal-bioceramic composite
US20110238155A1 (en) * 2007-09-28 2011-09-29 Abbott Cardiovascular Systems Inc. Stent formed from bioerodible metal-bioceramic composite
EP2397306A3 (en) * 2007-12-11 2013-03-27 Abbott Cardiovascular Systems Inc. Method of fabricating a stent from blow molded tubing
US8658082B2 (en) 2007-12-11 2014-02-25 Abbott Cardiovascular Systems Inc. Method of fabricating stents from blow molded tubing
EP2647484A1 (en) * 2007-12-11 2013-10-09 Abbott Cardiovascular Systems Inc. Method of fabricating a stent from blow molded tubing
US8414638B2 (en) * 2008-03-12 2013-04-09 Abbott Cardiovascular Systems Inc. Method for fabricating a polymer stent with break-away links for enhanced stent retenton
US20090234432A1 (en) * 2008-03-12 2009-09-17 Pacetti Stephen D Polymer Stent With Break-Away Links For Enhanced Stent Retention
US8852263B2 (en) * 2008-03-31 2014-10-07 Abbott Cardiovascular Systems Inc. Stent with nucleating agent
US20120277358A1 (en) * 2008-03-31 2012-11-01 Abbott Cardiovascular Systems Inc. Stent with nucleating agent
US10307511B2 (en) 2008-12-30 2019-06-04 Orthovita, Inc. Bioactive composites of polymer and glass and method for making same
US20100168798A1 (en) * 2008-12-30 2010-07-01 Clineff Theodore D Bioactive composites of polymer and glass and method for making same
US9662821B2 (en) 2008-12-30 2017-05-30 Orthovita, Inc. Bioactive composites of polymer and glass and method for making same
US9381277B2 (en) 2009-04-23 2016-07-05 Purac Biochem Bv Resorbable and biocompatible fibre glass compositions and their uses
KR101742017B1 (en) * 2009-04-23 2017-05-31 푸락 바이오켐 비.브이. Resorbable and biocompatible fibre glass compositions and their uses
US10933172B2 (en) 2009-07-24 2021-03-02 Warsaw Orthopedic, Inc. Implantable medical devices
US9399086B2 (en) 2009-07-24 2016-07-26 Warsaw Orthopedic, Inc Implantable medical devices
WO2012100651A1 (en) * 2011-01-27 2012-08-02 Dongguan Tiantianxiangshang Medical Technology Co., Ltd. Biodegradable stent formed with polymer-bioceramic nanoparticle composite and preparation method thereof
US20140039603A1 (en) * 2011-05-13 2014-02-06 Abbott Cardiovascular Systems Inc. Bioabsorbable scaffolds made from composites
US10213949B2 (en) 2011-08-11 2019-02-26 Abbott Cardiovascular Systems Inc. Controlling moisture in and plasticization of bioresorbable polymer for melt processing
CN105264014A (en) * 2013-05-16 2016-01-20 株式会社索夫塞拉 Biodegradable material
EP2998358A4 (en) * 2013-05-16 2017-02-08 SofSera Corporation Biodegradable material
US20150328373A1 (en) * 2014-05-19 2015-11-19 Abbott Cardiovascular Systems Inc. Additives To Increase Degradation Rate Of A Biodegradable Scaffolding And Methods Of Forming Same
US9480588B2 (en) 2014-08-15 2016-11-01 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9730819B2 (en) 2014-08-15 2017-08-15 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9855156B2 (en) 2014-08-15 2018-01-02 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US20180360628A1 (en) * 2014-08-15 2018-12-20 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9259339B1 (en) 2014-08-15 2016-02-16 Elixir Medical Corporation Biodegradable endoprostheses and methods of their fabrication
US9943426B2 (en) 2015-07-15 2018-04-17 Elixir Medical Corporation Uncaging stent
US10271976B2 (en) 2016-05-16 2019-04-30 Elixir Medical Corporation Uncaging stent
US10383750B1 (en) 2016-05-16 2019-08-20 Elixir Medical Corporation Uncaging stent
US10786374B2 (en) 2016-05-16 2020-09-29 Elixir Medical Corporation Uncaging stent
US10918505B2 (en) 2016-05-16 2021-02-16 Elixir Medical Corporation Uncaging stent
US10076431B2 (en) 2016-05-16 2018-09-18 Elixir Medical Corporation Uncaging stent
US11622872B2 (en) 2016-05-16 2023-04-11 Elixir Medical Corporation Uncaging stent
CN112545710A (en) * 2020-12-08 2021-03-26 中国人民解放军空军军医大学 Porous polyurethane support based on 3D prints

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US8512735B2 (en) 2013-08-20
US8309114B2 (en) 2012-11-13

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