Classifications

• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
  • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
  • A61L31/14—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
  • A61F2/90—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
  • A61F2/90—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure
  • A61F2/91—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
  • A61F2/90—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure
  • A61F2/91—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes
  • A61F2/915—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes with bands having a meander structure, adjacent bands being connected to each other
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/95—Instruments specially adapted for placement or removal of stents or stent-grafts
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
  • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
  • A61L31/04—Macromolecular materials
  • A61L31/06—Macromolecular materials obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
  • A61L31/00—Materials for other surgical articles, e.g. stents, stent-grafts, shunts, surgical drapes, guide wires, materials for adhesion prevention, occluding devices, surgical gloves, tissue fixation devices
  • A61L31/14—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
  • A61L31/148—Materials at least partially resorbable by the body
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
  • A61F2/90—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure
  • A61F2/91—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes
  • A61F2/915—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes with bands having a meander structure, adjacent bands being connected to each other
  • A61F2002/91533—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes with bands having a meander structure, adjacent bands being connected to each other characterised by the phase between adjacent bands
  • A61F2002/91541—Adjacent bands are arranged out of phase
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
  • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
  • A61F2/86—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure
  • A61F2/90—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure
  • A61F2/91—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes
  • A61F2/915—Stents in a form characterised by the wire-like elements; Stents in the form characterised by a net-like or mesh-like structure characterised by a net-like or mesh-like structure made from perforated sheet material or tubes, e.g. perforated by laser cuts or etched holes with bands having a meander structure, adjacent bands being connected to each other
  • A61F2002/9155—Adjacent bands being connected to each other
  • A61F2002/91558—Adjacent bands being connected to each other connected peak to peak

Definitions

• This invention relates to implantable medical devices, and particularly to bioresorbabie, biocompatible medical devices. Specifically, biocompatible, bioresorbabie stents useful in the treatment of strictures and preventing restenosis.
• PTCA percutaneous transluminal coronary angioplasty
• Stenosis of non-vascular tubular structures is often caused by inflammation, neoplasm and benign intimal hyperplasia.
• the obstruction can be surgically removed and the lumen repaired by anastomosis.
• the smaller transluminal spaces associated with ducts and vessels may also be repaired in this fashion; however, restenosis caused by intimal hyperplasia is common.
• dehiscence is also frequently associated with anastomosis requiring additional surgery, which can result in increased tissue damage, inflammation and scar tissue development leading to restenosis.
• BPH benign prostatic hypertrophy
• the most frequent cause is benign prostatic hypertrophy (BPH).
• BPH benign prostatic hypertrophy
• a number of therapeutic options are available for treating BPH. These include watchful waiting (no treatment), several drugs, a variety of so-called "less invasive” therapies, and transurethral resection of the prostate (TURP)--long considered the gold standard.
• Urethral strictures are also a significant cause of reduced urine flow rates.
• a urethral stricture is a circumferential band of fibrous scar tissue that progressively contracts and narrows the urethral lumen.
• Strictures of this type may be congenital or may result from urethral trauma or disease. Strictures were traditionally treated by dilation with sounds or bougies. More recently, balloon catheters became available for dilation. Surgical urethrotomy is currently the preferred treatment, but restenosis remains a significant problem.
• Permanent stents are used where long term structural support or restenosis prevention is required, or in cases where surgical removal of the implanted stent is impractical. Permanent stents are usually made from metals such as Phynox, 316 stainless steel, MP35N alloy, and superelastic Nitinol (nickel-titanium).
• Stents are also used as temporary devices to prevent closure of a recently opened urethra following minimally invasive procedures for BPH, which typically elicit post treatment edema and urethral obstruction. In these cases, the stent will typically not be covered with tissue (epithelialized) prior to removal.
• Temporary absorbable stents can be made from a wide range of synthetic biocompatible polymers depending on the physical qualities desired.
• Representative biocompatible polymers include polyanhydrides, polycaprolactone, polyglycolic acid, poly- L-lactic acid, poly-D-L-lactic acid and polyphosphate esters.
• Stents are designed to be deployed and expanded in different ways.
• a stent can be designed to self expand upon release from its delivery system, or it may require application of a radial force through the delivery system to expand the stent to the desired diameter.
• Self expanding stents are typically made of metal and are woven or wound like a spring. Synthetic polymer stents of this type are also known in the art.
• Self-expanding stents are compressed prior to insertion into the delivery device and released by the practitioner when correctly positioned within the stricture site. After release, the stent self expands to a predetermined diameter and is held in place by the expansion force or other physical features of the device. Stents that require mechanical expansion by the surgeon are commonly deployed by a balloon-type catheter. Once positioned within the stricture, the stent is expanded in situ to a size sufficient to fill the lumen and prevent restenosis. Various designs and other means of expansion have also been developed. One variation is described in Healy and Dorfman, U.S. Pat. No. 5,670,161. Healy and Dorfman disclose the use of a bio-compatible stent that is expanded by a thermo-mechanical process concomitant with deployment.
• urethral stents are composed of bio-compatible metals woven into a tubular mesh or wound into a continuous coil and are inserted endoscopically after opening the stricture by urethrotomy or sequential dilation.
• the stent is initially anchored in place through radial force as the stent exerts expansion pressure against the urethral wall.
• epithelial cells lining the urethra begin to grow through the stent's open weave between six and 12 weeks after insertion, thereby permanently securing the stent.
• U.S. Patent Nos. 5,670,161 a thermo-mechanically expanded biodegradable stent made from a co-polymer of L-lactide and ⁇ -caprolactone
• 5,085,629 a bioresorbabie urethral stent comprising a terpolymer of L-lactide, glycolide and ⁇ -caprolactone
• 5,160,341 a resorbable urethral stent made from polylactic acid or polyglycolic acid
• 5,441 ,515 a bio-erodible drug delivery stent and method with a drug release layer.
• bioresorbabie stents gradually hydrolyze in the body and stent fragments are then excreted, as in the case of urethral and bowel stents, or the nontoxic soluble degradation products may be absorbed and metabolized. Consequently, the use of bioresorbabie stents may ultimately eliminate the need for invasive removal procedures.
• advancements in polymeric, bio-resorbable stent design is still needed. Given, for example, there remains a need for bioresorbabie stents that provide enough radial strength to maintain luminal patency over a wide range of medical conditions and implantation sites. Furthermore, there is also a need to have bioresorbabie stents that have controlled degradation without total stent collapse and resulting obstruction. Moreover, there is a need for cost-effective biocompatible stents and processes for making stents that have differing functional lives.
• the present invention relates to implantable, bioresorbabie, biocompatible polymeric medical devices and methods for making same.
• the implantable, bioresorbabie, biocompatible polymeric medical devices of the present invention are intended for short to medium term in vivo use.
• the biocompatible, bioresorbabie medical devices of the present invention can be made from a variety of biocompatible polymeric compounds, their respective monomers, dimers, oligomers and blends thereof.
• the polymers used to make present invention include polyanhydrides, polycaprolactones, polyglycolic acids, poly- L-lactic acids, poly-D-L-lactic acids, and polyphosphate esters and their respective monomers, dimers, and oligomers.
• the polymeric materials of the present invention can be formed using techniques known to those having ordinary skill in the art of polymer chemistry and the material sciences. The polymers can be extruded into monofilaments, sheets or tubes and other configurations.
• the medical device is a biological stent, specifically a urethral stent.
• the medical device is a stent woven from a plurality of extruded polymeric monofilaments.
• the stent is extruded or injection molded as a tubular structure having fenestrations therein or provided with fenestrations thereafter using techniques known to those having ordinary skill in the art.
• Another embodiment of the present invention includes bioresorbabie stents having a radially self-expanding, tubular shaped member which may also expand and contract along its horizontal axis (axially retractable).
• the stent having first and second ends and a walled surface disposed between the first and second ends.
• the walled surface may include a plurality of substantially parallel pairs of monofilaments with the substantially parallel pairs of monofilaments woven in a helical shape.
• the stent is woven such that one-half of the substantially parallel pairs of monofilaments are wound clockwise in the longitudinal direction and one-half of the substantially parallel pairs of monofilaments are wound counterclockwise in the longitudinal direction. This results in a stent having an alternating, over-under plait of the oppositely wound pairs of monofilaments.
• Still another embodiment of the present invention may include a radially expandable, axially retractable bioresorbabie stent made from biocompatible, bioresorbabie polymers injection molded into a substantially tubular shaped device.
• the injection molded or extruded tubular shape device may have first and second ends with a walled structure disposed between the first and second ends and wherein the walled structure has fenestrations therein.
• the in vivo functional life of the stent is adjusted using methods comprising post-stent formation treatment steps selected from the group consisting of annealing, gamma irradiation and combinations thereof.
• the monomer content in the polymeric material is adjusted prior to stent formation using polymer extrusion pressure.
• the monomeric content is adjusted in the polymeric material using processes of blending polymeric and monomeric ingredients until a predetermined monomer is reached.
• Implantable polymeric medical devices include controlling the polymer's inherent morphology.
• a longer in vivo functional life is provided to a medical device by increasing the percentage of polymer having a crystalline morphology as opposed to an amorphous morphology.
• crystalline versus amorphous polymer morphology in the medical device is controlled using annealing temperatures and time.
• the present invention crystalline versus amorphous polymer morphology is controlled using monofilament draw ratio.
• methods for producing biocompatible, bioresorbabie stents having variable in vivo functional lives are provided wherein the ratio of monomer to high molecular weight polymeric sub-units in the polymer material used to form the polymeric stents is adjusted to achieve the desired in vivo functional life.
• FIG. 1 graphically depicts compression resistance of PLLA stents as a function of polymer viscosity over time in accordance with the teachings of the present invention
• FIG. 2 graphically depicts compression resistance of PLLA stents as a function of polymer viscosity over time in accordance with the teachings of the present invention
• FIG. 3 graphically depicts compression resistance of PLLA stents as a function of polymer viscosity over time in accordance with the teachings of the present invention
• FIG. 4 schematically depicts the manufacturing process for woven polymeric stents made in accordance with the teachings of the present invention
• FIG. 5 schematically depicts the manufacturing process for injected molded or extruded tubular polymeric stents made in accordance with the teachings of the present invention
• FIG. 6A is a side view of the bioresorbabie stent made in accordance with the teachings of the present invention.
• FIG. 6B is an end view of the bioresorbabie stent made in accordance with the teachings of the present invention.
• FIG. 6C is a perspective view of the bioresorbabie stent made in accordance with the teachings of the present invention.
• FIG. 7 is a side view of an alternate embodiment made in accordance with the teachings of the present invention.
• FIG. 8 is an enlarged view of a partial segment of the bioresorbabie stent made in accordance with the teachings of the present invention.
• FIG. 9 graphically depicts the bilateral self-expansion force of an alternate embodiment made in accordance with the teachings of the present invention versus UroLume® stents.
• FIG. 10 graphically depicts the bilateral compression resistance of one embodiment made in accordance with the teachings of the present invention versus UroLume® stents.
• FIG. 11 graphically depicts the radial self-expansion force by a Cuff Test of one embodiment made in accordance with the teachings of the present invention versus UroLume® stents.
• FIG. 12 graphically depicts the radial compression resistance by a Cuff Test of one embodiment made in accordance with the teachings of the present invention versus UroLume® stents.
• FIG. 13 graphically depicts the bilateral self-expansion force of one embodiment made in accordance with the teachings of the present invention as a function of in vitro aging time.
• FIG. 14 graphically depicts the bilateral compression resistance of one embodiment made in accordance with the teachings of the present invention as a function of in vitro aging time.
• FIG. 15 graphically depicts the radial compression resistance of an alternate embodiment made in accordance with the teachings of the present invention versus a UroLume® stent.
• FIG. 16 graphically depicts the radial self-expansion force of an alternate embodiment made in accordance with the teachings of the present invention versus a UroLume® stent.
• FIG. 17 graphically depicts the bilateral compression force versus calculated lumen area of bioresorbabie stents made in accordance with the teachings of the present invention.
• FIG. 18 graphically depicts the bilateral compression resistance as a function of time in vitro of various embodiments of bioresorbabie fenestrated tube stents made in accordance with the teachings of the present invention.
• FIG. 19 graphically depicts the bilateral self-expansion force as a function of time in vitro of various embodiments of bioresorbabie tube stents made in accordance with the teachings of the present invention.
• FIG. 20 schematically depicts the extrusion process used to make the monofilaments in accordance with the teachings of the present invention.
• Biocompatible A compound, composition of matter or device made therefrom that does not provoke more than a mild foreign body reaction in the host.
• Resorbable/Bioresorbable/Biodegradable A material that is broken-down in the body of the recipient into normal or non-toxic metabolic by-products. The resulting metabolic by-products are absorbed by the tissues and excreted from the body. A portion of the material may not be absorbed but rather be excreted in whole or in part by a physical action of the body such as peristalsis or urination without physical damage or toxic consequences to the recipient. Portions may also be resorbable.
• bioresorbabie, resorbable and biodegradable may be used interchangeably when describing certain embodiments of the present invention. Unless specifically contradicted by the text, no distinction is to be made between these terms when used in conjunction with urethral stents.
• Implantable/lmplant Mechanically or surgically placed into the body of the recipient.
• Polymeric sub-units A monomer, dimer, or oligomer of the basic polymer chain.
• Polymeric Ingredients Polymeric sub-units.
• Polymeric composition A polymeric material composed of at least one polymeric ingredient of at least one type of polymer.
• the stents of the present invention are intended for in vivo use ranging from 1-3 months for "short-term” applications and 3-6 months for "medium-term” applications.
• In vivo functional life The point at which a polymeric stent has less than 50% of its initial compression resistance as measured in Newtons.
• Draw Ratio This is the ratio of the roller speed at the last godet station to that of roller speed at the first godet station as depicted in Figure 20.
• High molecular weight polymer A polymer having an inherent viscosity greater than 4.5 dl/g.
• Low molecular weight polymer A polymer having an inherent viscosity less than 4.5 dl/g.
• the present invention relates to polymeric medical devices that are implanted into the body of a patient in need thereof.
• the medical devices of the present invention are designed to be biocompatible and bioresorbabie. Biocompatibility is required to enable the medical device to remain in the patient for a sufficient time to provide its intended benefit without provoking an adverse host response. Biocompatibility is achieved by selected materials that are relatively inert, or that are recognized by the host as "self.” For example, many metals are chemically and biologically inert. Examples include stainless steel, titanium nickel alloys and mixtures thereof. Inert materials may also include polymers, or "plastics" that are made from a wide variety of monomeric sub-units.
• metal alloys include, for example, skull plates, artificial joints, supports for damaged bones and bone screws.
• metal alloys may be too bulky, rigid or subject to chemical attack and encrustation.
• medical implants made from metal alloys must either be permanently implanted, or surgically removed. There are many applications where temporary applications are preferred. In these cases, medical devices made from bioresorbabie materials that will not require post implantation surgical removal may be preferred.
• Bioresorbabie medical polymers were first used in the 1970s when resorbable sutures made from Dexon ® where introduced.
• Dexon ® is poly-glycolic acid polymer (a poly-alpha-hydroxy acid) composed of glycolic acid sub-units.
• Poly-glycolic acid (PGA) polymers are degraded in the body by hydrolysis into oligomers that in turn are broken down into glycolic acid monomers. These glycolic acid monomers are ultimately broken-down into pyruvic acid and finally metabolized into carbon dioxide and water. Since the successful introduction of Dextron ® many other biocompatible, bioresorbabie polymers have been used to make medical devices. There are numerous factors that must be considered when selecting a polymer material for use as a medical implant.
• Structural strength of the implant is just a few of the considerations.
• a wide variety of surgical procedures and applications are contemplated herein.
• the present invention is believed particularly suitable for use in conjunction with surgical procedures for treating the prostate or the lower urinary tract.
• a patient undergoing brachytherapy may have a short term stent implanted to resist blockage of the urinary tract due to swelling of the prostate.
• a procedure for treating the prostate e.g. a Trans-Urethral Resection of the Prostate [TURP], microwave therapy, RF treatment, or the like
• the stent may be used in conjunction with a treatment for a urethral stricture to help resist any tendency for the tissue to grow together or occlude the urinary tract.
• the urethral stents of the present invention are intended for short to medium term applications. Therefore, in one embodiment of the present invention, the stents are made from a polymeric composition designed to be resorbed within in a specified time period. However, in order to provide the recipient its intended benefit, the stent must retain sufficient structural integrity to maintain a minimum compression resistance over its intended in vivo life span. Therefore, resorption must occur gradually. Consequently, the present inventors have developed stents having specific polymer compositions and structural features that fulfill the combined objectives of short to medium term structural strength with bioresorbability.
• Physical properties of polymers are influenced by the size of the molecules and by the nature of the primary and secondary bond forces.
• the type and size of monomers, polymer sub-units, overall polymer viscosity and polymer morphology influence these properties.
• the present inventors have determined that a polymer's in vivo bioresorption rate and structural strength are a function of these physical properties.
• Monomer content can significantly affect in vivo functional life. Specifically, increasing the monomer content in polymeric medical devices made from polymers having high initial molecular weights significantly shortens in vivo functional life.
• polymer morphology also contributes to bioresorption rates. While not as significant as the monomer percentage in the final polymeric composition, the present inventors have demonstrated that increasing the device's amorphous domains relative to its crystalline domains can decrease the polymer's in vivo functional life.
• the synthetic polymers of the present invention are produced by a process governed by random events. As a result, the chain lengths of individual polymer sub- units vary. Consequently, a particular polymeric material cannot be characterized by a single molecular weight. Instead, a statistical average of all of the polymeric sub- units is used to denote molecular weight.
• the molecular weight of polymers can be expressed in different ways including number average, weight average and viscosity average. Number average is the sum of all molecular weights of the individual molecules present divided by their total number. In weight averages each polymeric sub-unit contributes according to the ratio of its particular molecular weight to the total.
• viscosity average is best suited for linear polymers such as those used in the foregoing examples. Therefore, for these reasons, the viscosity average method will be used throughout this specification to determine and denote the molecular weights.
• the present inventors have determined that a polymer's monomer content (measured as a percentage of total polymeric subunits in a polymeric medical device) is directly related to polymer stability under hydrolytic conditions. Hydrolytic stability in turn affects bioresorption rates and hence a medical device's in vivo functional life.
• the present inventors have also determined that hydrolytic stability is also affected by the polymer's morphology.
• the present inventors have determined that polymer morphology is affected by physical factors such as initial draw ratio, annealing temperature, annealing time and the extent of contraction allowed during annealing.
• Example I details methods used to determine polymer inherent viscosity.
• Example 2 provides methods for determining polymer monomer content using nuclear magnetic resonance (NMR) testing.
• Example 3 teaches a polymer extrusion process.
• Example 4 details the method used to test bilateral compression resistance of stents made in accordance with the teaching of the present invention.
• Example 5 describes the methods used to simulate the in vivo hydrolytic environment. Stents incubated under the conditions and for the times described in Example 5 were used to assess polymer performance as a function of time under physiological conditions.
• Linear polymer solution viscosity relates to average molecular weight and can be used to designate polymer size.
• Capillary efflux time ( ) of a polymer dissolved in an appropriate solvent is measured at constant temperature and compared with the efflux time for pure solvent (to) at the same temperature. These values are then used to calculate polymer inherent viscosity. While this example uses poly-L-lactic acid (PLLA) polymer, this is not intended as a limitation. The following example can be used to determine the inherent viscosity for many polymers, specifically linear polymers.
• PLLA poly-L-lactic acid
• Inherent Viscosity Measurement a) After thoroughly rinsing and drying (aspirate) the viscometer, measure 10ml of chloroform in a volumetric pipet. Dispense into the clean viscometer and close the lid. Click on Viscometer (stand) icon of choice at the screen and fill in the sample ID, lot number, operator name, etc. Choose kinematic viscosity and click on start to run the standard chloroform sample, automatically.
• a polymer is dissolved in an appropriate solvent and examined by NMR to determine its structure. Resonance areas are measured to determine the percent composition of the polymer, the residual monomer and any significant impurities present.
• Polylactide can be analyzed in deuterochloroform (CDCI 3 ).
• Spectra may be run at any temperature between 20°C and 45°C, typically at 35°C. Resonance positions will shift slightly with temperature changes. i. Spectra are run under quantitative conditions:
• EXAMPLE 3 Extrusion Process Polymer granules are loaded in the hopper 201 of the extruder 202.
• the extruder screw 203 in the heated barrel melts the polymer and delivers it to metering pump (not shown) under pressure.
• the metering pump pushes the melt through a spin head 204.
• a spin head consists of a 'screen pack' to filter the melt and a spinneret die.
• Molten monofilament strands are quenched in a water bath 205. The quenched strands pass over the rollers 206 of the first godet station 207.
• the speed of the rollers 206 of godet station 207 is adjusted to match the flow rate through the spin head 204.
• the strands then pass through a drawing oven 208 and subsequently over the rollers 209 of the second godet station 210.
• the speed of the rollers 209 at godet station 210 is faster than roller 206 at the first godet station 207 to apply initial draw to the monofilament strands.
• the strands then pass through the next set of drawing oven and godet station
• the rollers at this godet station rotate at even higher speed to apply additional draw to the strands.
• the strands are next collected over spools on traverse winder 211.
• Bi-lateral compression/relaxation (BLCR) testing used to determine the compression resistance and self expansion force of polymeric stents made in accordance with the teachings of the present invention.
• Supplies, apparatus and reagents a) Instron, Model 5565 with Merlin Test Profiler software b) Instron load cell, 200 lb. c) Bi-lateral Compression/Relaxation test fixture d) Caliper (mm., calibrated to two decimal places)
• Profiler Click on "Profiler” and verify the following, then exit Profiler: Tensile Extension, Relative ramp, 1 Ramp, #1 , Delta at 10.50 mm, and Rate at -5.0 mm/min. Click on the right pointing arrow at the top of the screen to verify subsequent blocks or ramp parameters as follows: Tensile Extension, Hold, 2 Hold, #2, Duration: 1 minute. Tensile extension, Relative ramp, 3 Ramp, #3, Delta at 10.50 mm, and Rate at 2.0 mm/min.
• sample ID's i.e., 1.5, 2.0, 2.5, or 3.0 cm
• nominal stent length i.e., 1.5, 2.0, 2.5, or 3.0 cm
• TTD in mm i.e., 1.5, 2.0, 2.5, or 3.0 cm
• BLCRstnt BLCRstnt
• the test objective is to characterize the compression resistance and the self- expansion (S-E) force of braided PLLA stents.
• the raw data of crosshead displacement versus force must be treated to obtain the platen gap versus force data for the stent.
• the data characterize the two cycles consisting of the following 3 sequential steps: a) In the first step the stent was compressed to a controlled outside diameter
• the data for platen gap versus force from each sample are to be treated to determine two parameters used to describe the stent's mechanical properties.
• the two parameters are S-E force in the first cycle and compression resistance in the second cycle.
• the self-expansion force and compression resistance measured at 10 mm platen gap are reported as representative measures of the respective stent properties.
• All stents will be 3.0 cm long at 14 mm OD.
• dissolved air will be removed using the following procedure: Place bottles with stents and PBS in the vacuum oven. Remove all bottle caps. Close the vacuum chamber and gradually reduce chamber pressure to approximately 0.090-0.095 MPa. As the pressure declines, watch for growth of air bubbles on the stents.
• the raw data of crosshead displacement versus force will be converted to platen-gap versus force for each stage of the BLCR test.
• Figures 1-3 plot the mean values for the stent samples used in Example 5 and tested in accordance with the teachings of Example 4.
• Figures 1-3 demonstrate a direct correlation between increasing levels of gamma radiation used to treat stent samples and a reduction in initial inherent viscosity and compressive strength.
• FIGs. 1-3 also demonstrate that as the stents are maintained under simulated in vivo conditions, compressive strength diminishes over time in a direct relationship to reduction in overall inherent viscosity.
• the present inventors have determined that there are numerous factors that influence the size distribution of polymeric molecules in a polymeric composition. Specifically, the present inventors have identified several physical factors that can be used to control the monomer content in the polymeric medical devices of the present invention.
• polymers used to fabricate implantable medical devices can be derived in a number of different ways.
• a polymer is selected having monomer content within a predetermined range.
• the polymer is then pelletized, milled and extruded into the appropriate configuration.
• the polymer is a blend of polymer compositions selected from a number of different molecular weights. The mixture is then blended, pelletized and then extruded.
• the term "predetermined range" is defined as a value selected based on the teachings of the present invention that will result in the medical device having the functional qualities desired. For example, and not intended as a limitation, a urethral stent having a compression resistance in Newtons (N) of 7.0 with a useful in vivo life span of five weeks is desired (the useful in vivo life span in the present example is defined as the time at which the stent will have a minimum compression resistance in N of ⁇ 3.5). Based on the teachings of the present invention, it is determined that a polymer stent made using a high molecular weight (high inherent viscosity) polymer and having a monomer content 1.2 weight percent (wt. %) to 2.0 wt. % would be required. This range in monomer wt. % would be the "predetermined range.”
• a medical device's useful in vivo life span is principally determined by the time it takes to lose 50% or more if its initial structural strength.
• polymeric urethral stents will be the medical device and "structural strength" will be measured by the stent's ability to maintain lumen patency for a specific period (compression resistance as measured by the BLCR test of Example 4). Therefore, in the discussion that follows a stent's structural strength will be its compression resistance measured in Newtons. Therefore, a stent's in vivo functional life is defined as the amount of time an implanted stent will retain at least 50% of its initial compression resistance once exposed to a hydrolytic (in vivo) environment.
• the stents of the present invention are intended for short to medium term use.
• the average in vivo functional life for the bioresorbabie stents of the present invention range from approximately 1-3 months for "short-term” applications and 3-6 months for "medium-term” applications.
• a polymeric stent's structural strength diminishes in vivo as a result of hydrolytic activities.
• bioresorbabie polymers possess regions within the polymer matrix that are subject to attach by water under physiological conditions. As the polymer matrix undergoes hydrolytic attack, it is broken down into smaller polymeric subunits that are eventually metabolized at the cellular level through the citric acid cycle into water, carbon dioxide and energy. Thus the polymer matrix is weaken by the combined processes of fragmentation and net polymer viscosity reduction.
• the present inventors have ascertained that there are two fundamental polymer properties that can be modulated during the manufacturing process to control the rate of hydrolytic attack.
• Table 1 depicts the in vitro functional lives of woven urethral stents made from extruded poly-L-lactic acid (PLLA) monofilaments in accordance with the teachings of the present invention.
• the stents were subjected to in vitro stability testing as detailed in Examples 4 and 5 above.
• Table 1 demonstrates that stents fabricated using polymers having an initial inherent viscosity of 8.0 dl/g or above lose compression resistance more rapidly as the monofilament monomer content increases (providing the annealing conditions are constant).
• Polymer morphology also affects polymeric stent in vivo functional life.
• Polymeric compositions may be primarily crystalline, amorphous or a combination thereof.
• Crystalline polymers are generally composed of symmetrical polymer chains that permit the individual polymer molecules to stretch out straight and align themselves with each other. It is well known in the art of polymer chemistry that most polymers do not fully stretch out, but rather are composed of molecules that fold back on themselves forming structures known as lamellae. This is particularly true for high molecular weight polymeric subunits that have a great deal of intramolecular symmetry such as high viscosity PLLA.
• the lamellae form neatly packed polymer crystals that are tightly packed and resist hydrolytic attack because water does not easily penetrate the hydrophobic regions of the polymer molecule.
• most crystalline polymers may have amorphous regions formed by portions of the polymer chain that do not readily align themselves with the lamellae. The amorphous regions are not susceptible to hydrolytic attack. Therefore, the more amorphous regions in a polymer, the faster it may be degraded in a hydrolytic environment.
• the dominant factor affecting in vivo hydrolytic degradation is the percent monomer content and the molecular weight (inherent viscosity) of the pre-processed polymer component.
• the present inventors have ascertained that short and medium term in vivo functional lives are most effectively controlled using a high molecular weight polymer (8.0 or greater dl/g) as the starting material and increasing monomer content in the final polymer composition.
• monomer content in the final polymer composition e.g. a monofilament or stent
• pre-formation phase There is essentially phase in the manufacturing of the present medical devices wherein the polymer composition's monomer content may be altered to achieve a predetermined range.
• pre-formation phase includes, but may not be limited to, dry blending (10/20), extruding polymer rods (11/21 ), pelletizing extruded rods (12/22), drying pellets (13/23), extruding coarse monofilaments (14), melting pellets in injection molder (24), quenching (15), injection molding (25), drawing the final monofilament (16), and unmolding (26).
• the medical devices of the present invention can therefore be fabricated to have a final monomer content within a predetermined range.
• Pre-formation steps also include determining the selected polymer's inherent monomer content using methods known to those skilled in the art of polymer chemistry.
• monomer content is determined using NMR techniques.
• the monomer content of the starting material is compared to the predetermined monomer content for the monofilament or finished stent having the in vivo functional life desired (the predetermined range). If the monomer content is below the predetermined percentage, monomer content is adjusted using one or more pre- formation techniques.
• monomer content is adjusted by adding monomer to the polymer prior to the blending or extrusion processes. In another embodiment polymer extrusion conditions is used to increase monomer content in the polymer composition.
• a bioresorbabie stent having an initial compression resistance of 6 N and a useful in vivo life of eight weeks. Consequently, if the initial polymer selected to make this particular stent has an initial inherent viscosity of 8.0 dl/g, then it can be determined from Table 1 that the monomer content of the pre-annealed, pre-irradiated polymer must be below 1.4%. Preferably the monomer content is between approximately 1.1 and 1.31 %.
• the polymer composition used in this non-limiting example may be prepared by blending high molecular weight PLLA preparations to obtain the predetermined monomer range or a high molecular weight PPLA may be extruded at a pressure such that the predetermined amount of PLLA monomer is formed in the polymer composition prior to completing stent fabrication.
• a combination of methods may be used to achieve the predetermined monomer content.
• stents made in accordance with the teachings of the present invention may be treated after fabrication in order to achieve desired bioresorbability rates.
• these post fabrication processes include, but are not limited to, exposing the finished stent to different doses of gamma irradiation from a Cobalt 60 source and/or annealing the stent at different temperatures and for different times.
• monomer content can be monitored through out the manufacturing process to verify that the predetermined monomer content is achieved.
• the exact monomer content can be achieved by using NMR, or other techniques, to monitor the stent's monomer content during manufacturing (in process testing).
• NMR or other techniques
• the amount of energy used 35 kGy to 50 kGy respectively is greater than that used for sterilizing medical devices.
• higher doses of radiation are used in the present invention randomly decrease the molecular weight of the high molecular weight polymeric sub-units.
• the stents described herein have also been subjected to annealing in order to achieve the initial compression resistance desired.
• the stent's physical configuration will dramatically affect its overall structural integrity and, thus, in vivo life span.
• Stents woven from monofilaments have different physical qualities than stents made from solid extruded tubes having fenestrations cut therein.
• the monofilament diameter as well as the number of stands and braiding pattern have a significant impact on stent strength and, thus, in vivo life.
• Tensile strength is defined as the force per unit cross- sectional area at the breaking point. It is the amount of force, usually expressed in pounds per square inch (psi), that a substrate can withstand before it breaks, or fractures.
• the tensile modulus, expressed in psi is the force required to achieve one unit of strain, which is an expression of a substrate's stiffness, or resistance to stretching and relates directly to the stent's performance.
• the filament possesses a tensile strength in the range from about 40,000 psi to about 120,000 psi with an optimum tensile strength for the filament 30 of approximately between 60,000 to 120,000 psi.
• the tensile strength for the fenestrated stent 23 is from about 8,000 psi to about 12,000 psi with an optimum of about 8,700 psi to about 11 ,600 psi.
• the tensile modulus of polymer blends in both embodiments ranges between approximately 400,000 psi to about 2,000,000 psi.
• the optimum range for a stent application in accordance with the present invention is between approximately 700,000 psi to approximately 1 ,200,000 psi for the woven embodiment and approximately 400,000 psi to 800,000 psi for the fenestrated embodiment.
• a single PLLA formulation having a predetermined inherent viscosity may be used alone, or it may be blended with one or more PLLA compositions having different inherent viscosities and/or differing amounts of PLLA monomer.
• the exact number of steps used to make all possible embodiments of the present invention will vary depending upon the whether polymer blends are used, or whether a single polymer having a predetermined inherent viscosity is used and how much and/or if any additional monomer is added.
• the manufacturing process will begin with dry blending under an inert atmosphere (10 in FIG. 4 or 20 in FIG. 5).
• an inert atmosphere 10 in FIG. 4 or 20 in FIG. 5
• the process may begin by extruding polymer rods (11 in FIG. 4 or 21 in FIG 5) or by adding pellets (13 in FIG. 4 of 23 in FIG. 5) directly to either an extruder (14 in FIG 4) or injection molder (24 in FIG. 5).
• Figure 4 depicts the basic steps for making one embodiment of the present invention.
• one or more polymer compositions are selected such that the final monofilament will have monomer content within a predetermined range.
• the polymer composition(s) is dry blended 10 under an inert atmosphere, then extruded in rod form 11.
• the polymer rod is pelletized 12 then dried 13.
• the dried polymer pellets are then extruded 14 forming a coarse monofilament that is quenched 15.
• the extruded, quenched, crude monofilament is then drawn into a final monofilament 16 with an average diameter from approximately 0.145 mm to 0.6 mm, preferably between approximately 0.35 mm and 0.45 mm.
• Approximately 10 to approximately 50 of the final monofilaments 16 are then woven 17 in a plaited fashion with a braid angle 46 (FIG. 6A), from about 100 to 150 degrees on a braid mandrel of about 3 mm to about 30 mm in diameter.
• the plaited stent 30 (FIG. 6A) is then removed from the braid mandrel and disposed onto an annealing mandrel having an outer diameter of equal to or less than the braid mandrel diameter and annealed 18 at a temperature between about the polymer-glass transition temperature and the melting temperature of the polymer blend for a time period between about five minutes and about 18 hours in air, an inert atmosphere or under vacuum.
• the stent 30 (FIG. 6A) is then allowed to cool and is then cut 19.
• a first step 20 may include blending one or more polymers or a single polymer using multiple inherent viscosities. The blending is done in an inert atmosphere or under vacuum. The polymer is extruded in rod form 21 , quenched 21 , and then pelletized 22. Typically, the polymer pellets are dried 23, then melted in the barrel of an injection molding machine 24 and then injected into a mold under pressure where it is allowed to cool and solidify 25. The stent is then removed from the mold 26. The stent tube may, or may not, be molded with fenestrations in the stent tube. In one embodiment of the fenestrated stent 50 (FIG.
• the tube blank is injection molded or extruded, preferably injection molded, without fenestrations.
• fenestrations are cut into the tube using die-cutting, machining or laser cutting, preferably laser cutting 27.
• the resulting fenestrations, or windows may assume any shape, which does not adversely affect the compression and self- expansion characteristics of the final stent.
• the stent is then disposed on an annealing mandrel 28 having an outer diameter of equal to or less than the inner diameter of the stent and annealed at a temperature between about the polymer-glass transition temperature and the melting temperature of the polymer blend for a time period between about five minutes and 18 hours in air, an inert atmosphere or under vacuum 28.
• the stent 50 (FIG. 7) is allowed to cool 29 and then cut as required 30.
• FIGs. 6A-6C depict a bioresorbabie, self-expanding stent 30.
• Figures 6A-6C show the bioresorbabie stent 30 comprising a cylindrical sleeve having a first end 38 and a second end 40.
• a plurality of monofilaments 32 which are positioned substantially parallel and helically wound about the longitudinal axis 34 of the stent 30 to form a latticed network 35.
• the latticed network 36 forms the wall 42 of the bioresorbabie stent. As shown in Figs.
• the monofilaments 32 are braided in an alternating under-two-over-two pattern forming the latticed network.
• the braid-crossing angle 46 is the obtuse angle between any two monofilaments 32 at a point of intersection.
• thirty to forty-eight monofilaments may be braided to form the bioresorbabie stent 30; preferably forty monofilaments are braided to form the bioresorbabie stent.
• the present invention also contemplates braiding patterns such as, but not limited to, under-one-over-one, under-one-over-two, under-one-over-three, under-two-over-three, under-three-over-three, and the like.
• Figs 6A-6C uneven openings result as shown in Figs 6A-6C. That is, the openings in the latticed network are not uniform.
• uniform openings may be provided in a bioresorbabie stent by manufacturing the stent on a braiding device with the appropriate number of evenly spaced carriers. For example, a thirty- strand stent may be formed on a 30 carrier braiding device. Uniform openings may also be achieved by pairing strands in a 48-strand stent with the under-two-over-two braid pattern.
• FIG. 8 is an enlarged view showing the under-two-over-two braiding pattern of the bioresorbabie stents 30, 30' of the present invention. Furthermore, Fig. 8 illustrates a bioresorbabie stent 30' having a single strand shift.
• a single strand shift is defined as adjacent monofilaments 32', 33' having a different braiding pattern. For instance, a monofilament 32' will have an under-two-over-two braiding pattern and the adjacent monofilament 33' will have an under-two-over-two braiding pattern offset by one monofilament. Stated differently, any adjacent monofilaments will not go "under and over" the same monofilaments.
• Figs. 6A-6C also show openings 44 between the individual monofilaments 32 that comprise the latticed network 35 of the stent 30. Providing spaces throughout the latticed network 35 of the stent 30 allows for sufficient tissue in-growth between the monofilaments of the latticed network thereby fixing the stent in position and minimizing the likelihood of stent migration or dislodgment.
• bioresorbabie stents having openings of different sizes are also contemplated in the present invention provided that suitable self-expansion forces and compression resistance are achieved.
• Figs. 9-10 graphically depict the bilateral self-expansion forces and compression resistance forces of one embodiment of the present invention versus UroLume® stents.
• UroLume® is the trademark for a metallic stent marketed by American Medical Systems, Inc., the assignee of the current application.
• Figs. 9-10 graphically depict the bilateral self-expansion forces and compression resistance forces of one embodiment of the present invention versus UroLume® stents.
• UroLume® is the trademark for a metallic stent marketed by American Medical Systems, Inc., the assignee of the current application.
• Figs. 9-10 graphically depict the bilateral self-expansion forces and compression resistance forces of one embodiment of the present invention versus UroLume® stents.
• UroLume® is the trademark for a metallic stent marketed by American Medical Systems, Inc., the assignee of the current application.
• Figs. 9-10 graphically depict the bilateral self-ex
• 9-10 graphically compare bioresorbabie stents having 40 poly-L-lactic acid monofilaments braided in an under-two-over-two pattern and treated at various gamma irradiation doses (35 kGy, 50kGy, and 65kGy) versus UroLume® stents having braid-crossing angles of 118° and 145°.
• the stent samples were subjected to a bilateral compression-relaxation test using an Instron test machine.
• the stents were compressed bilaterally between two smooth platens of a Delrin fixture from a resting state to a platen gap of 7 mm.
• the platen gap range of 7 mm to 15 mm corresponds to the stent diameter in a compressed state (7 mm) and an expanded state (15mm).
• the stents were held for a set hold-time of approximately 1 minute, and the stents were allowed to relax.
• the stents were subjected to two cycles of compression, hold, and relaxation.
• the force exerted by the stent during the relaxation stage of the first cycle was recorded as the self-expansion force.
• the force applied to compress the stent in the second cycle was recorded as the compression resistance of the stent.
• Fig. 9 illustrates that the bioresorbabie stents of the present invention have better bilateral self-expansion forces as compared to the UroLume® stents over a platen gap range of 7 mm to 15 mm.
• a bioresorbabie stent exposed to 35 kGy dose of gamma irradiation exerts a bilateral self-expansion force of approximately 9 N while UroLume® stents having braid- crossing angles of 118° or 145° exert self-expansion forces of 3N and approximately 5 N, respectively.
• Figs. 11-12 also show similar results when the stents of the present invention and UroLume® stents were subjected to a Cuff test.
• the Cuff test was conducted on an Instron test machine using a test fixture and a Mylar® collar.
• the test fixture consists of a pair of freely rotating rollers separated by a 1-mm gap, and the Mylar® collar is a laminated film of Mylar® and aluminum foil.
• a 30-mm long stent segment was wrapped in a 25-mm wide collar and the two ends of the collar were passed together through the rollers of the test fixture. A pulling force was applied to the collar ends, which radially compressed the stent against the rollers.
• the stent samples were compressed from their resting diameter to a predetermined diameter (typically 7-mm).
• the stent samples were compressed and held at the predetermined diameter for approximately one minute, and then they were allowed to relax.
• the stents were subjected to two cycles of compression, hold and relaxation.
• the force exerted by the stent during the relaxation stage of the first cycle was recorded as the self-expansion force.
• the force applied to compress the stent in the second cycle was recorded as the compression resistance of the stent.
• the bioresorbabie stents of the present invention demonstrated greater radial self-expansion forces over the whole range of constrained stent diameters from 7mm to 15 mm as compared to the UroLume® stents.
• the bioresorbabie stents displayed approximately 9 N to 11 N of radial self-expansion force at a constrained stent diameter of 7 mm as compared to 3 N and 5 N at 7 mm of radial self-expansion force for the UroLume® stents, as shown in Fig. 10.
• the superior results are also illustrated by the graphical data in Fig. 11.
• the graphical data set forth in Figs. 9-11 illustrate that the bioresorbabie stents having an under-two-over-two braided pattern have superior radial self-expanding forces and compression resistance forces as compared to UroLume® metallic stents.
• the bioresorbabie stents of the present invention are also controllably biodegradable, which eliminates the need for complicated or invasive stent removal procedures. That is, once an implanted stent has served its intended function, the stent is controllably degraded and naturally eliminated by the human body.
• the bioresorbabie, self-expanding stents are manufactured by providing a plurality of monofilaments and braiding these monofilaments in an under-two-over two pattern to form a latticed network as shown in Fig. 6 and Fig. 8.
• the latticed network of the bioresorbabie stents comprises thirty to forty-eight monofilaments.
• the latticed network is formed by winding the monofilaments about a mandrel. Approximately half of the monofilaments are wound around the mandrel in a clockwise direction while the other half of the monofilaments are wound in a counter-clockwise direction.
• the angle between the two filaments at the point where they intersect is defined as the braid-crossing angle 46 as shown in Fig. 6. It is contemplated that the monofilaments intersect at a braid-crossing angle between 100° to 150°.
• the bioresorbabie stents comprise monofilaments having an as-braided braid-crossing angle of 110°. Those skilled in the art will appreciate that/other braid-crossing angles may be selected to achieve different self-expansion forces or compression resistance.
• the bioresorbabie stents then undergo an annealing process.
• the annealing process includes placing the bioresorbabie stents on a mandrel, axially compressing the stents by 30% to 60%, heating the stents to the glass transition temperature of the biocompatible polymer for a predetermined period of time, and allowing the stents to be controllably cooled.
• the annealing process relieves internal stresses and instabilities of the monofilaments that result from the production of the bioresorbabie stents.
• the bioresorbabie stents are heated to approximately 90°C for a length of time between about one and about eight hours, preferably four hours, in an inert atmosphere.
• the inert atmosphere may be comprised of a high vacuum or nitrogen gas.
• argon, or helium a gas that is also contemplated including, but not limited to, argon, or helium.
• the bioresorbabie stents are then controllably cooled to room temperature. Each stent is then cut to desired size for its intended application.
• the stents are exposed to Co 60 gamma irradiation to fine tune the in vivo functional life of the bioresorbabie stents. Exposure to gamma irradiation causes molecular degradation of the polymers that comprise the bioresorbabie stents; however, the gamma irradiation does not affect the overall morphology of the polymers.
• the monofilaments that comprise the bioresorbabie stent contract resulting in a different final braid-crossing angle.
• the contraction of the monofilaments that comprise the braided stent is important in achieving the compression resistance and self-expansion forces for the stents of the present invention.
• the final post-annealing braid angle ranges from approximately 125° to 150°, and more particularly a final braid angle ranging from approximately 130° to 145°.
• the final post- annealing braid angle is dependent upon the desired properties and stent length. For instance, a 1.5 cm long stent would require a final post-annealing braid angle ranging from approximately 139° to 145° whereas a lesser braiding angle might be adequate for a longer stent.
• the in vivo functional life of the bioresorbabie stents is related to the temperature and duration of the annealing process and the dosage of gamma irradiation. Accordingly, the functional lifetime of the stents can be controlled and/or adjusted by manipulating the annealing conditions during the manufacturing process.
• the annealing conditions of 90°C for a length of time between about one to about eight hours, preferably four hours, in an inert atmosphere followed by 50 kGy dose of gamma irradiation provides bioresorbabie stents having approximately a two week functional life and substantial stent degradation by approximately the fourth week of in vivo implantation.
• the bioresorbabie stents may be annealed at a temperature higher than 110°C for at least eight hours to achieve an in vivo functional life between three to six months.
• the bioresorbabie stents are typically annealed at 1 0°C for approximately eighteen hours to achieve an in vivo functional life between three to six months.
• the annealing parameters may be adjusted for shorter or longer in vivo functional lives.
• Figs. 13-14 graphically illustrate the mechanical strengths of the bioresorbabie stents of the present invention as a function of in vitro aging time.
• the in vitro study parameters were designed to mimic in vivo functional life. Accordingly, the stents were aged in a phosphate buffered saline (pH 7.3) at 37°C, and samples were then tested in a bilateral compression/relaxation test at each corresponding aging period.
• Figs. 13-14 show the changes in the self-expansion force and bilateral compression resistance of the bioresorbabie stents over a six week period of time. For instance, as shown in Figs. 13-14, the stents exposed to 35 kGy and 50 kGy doses of gamma irradiation retained >70% of their initial mechanical strength for two weeks, but a substantial degradation in mechanical strength had occurred by the fourth week.
• Fig. 7 illustrates a second embodiment of the present invention.
• the second embodiment of the present invention is similar to the laser cut stent as disclosed in U. S. Patent Number 5,356,423, the entire contents are herein incorporated by reference.
• the bioresorbabie stent 50 is comprised of a tubular sheath 52 having a first end 54 and a second end 56.
• a walled surface 58 having a plurality of fenestrations 60 spaced throughout the walled surface 58 is shown in Fig. 7.
• the walled surface 58 is contemplated to have a thickness of 0.025" to 0.030", preferably 0.030".
• the fenestrations 60 are shaped in such a manner to maximize the number of openings for tissue in-growth while maintaining the predetermined self-expansion and compression resistance forces of the bioresorbabie stent.
• the bioresorbabie stents are formed by the following process.
• Bioresorbabie, biocompatible polymers are injection molded or extruded into a tubular sheath.
• the polymers may be selected from any known bioresorbabie polymers including, but not limited to, polyanhydrides, polycaprolactones, polyglycolic acids, poly-L-lactic acids, poly-D-L-lactic acids, polydioxanone, and polyphosphate esters.
• polydioxanone is used to form the tubular sheath.
• the tubular sheath may be injection molded with or without fenestrations.
• the tubular sheath is injection molded without fenestrations.
• the fenestrations are introduced into the tubular sheaths by cutting processes including, but not limited to, laser cutting and machining.
• the bioresorbabie stents then undergo an annealing process.
• the annealing process includes heating the stents to or above the glass transition temperature of the biocompatible polymer for a predetermined period of time, and allowing the stents to cool slowly.
• the annealing process relieves internal stresses and instabilities that result from the production of the bioresorbabie stents of the present invention.
• Bioresorbabie stents made from polydioxanone are heated to a temperature of approximately 75°C for between about one and six hours, preferably three hours, in an inert atmosphere of high vacuum or nitrogen gas and controllably cooled for approximately twelve hours.
• inert atmospheres having low moisture content are also contemplated including, but not limited to, argon, or helium.
• Figs. 15-16 illustrates the mechanical properties of the bioresorbabie stent 50.
• Figs. 15-16 graphically depict the radial compression resistance and self-expansion forces of two embodiments of the bioresorbabie stent 50 having different fenestration designs and wall thickness versus a 145° Urolume® stent.
• the stent samples were subjected to a Suture test using an Instron test machine.
• the Suture test is similar to the Cuff test with the exception that a suture, rather than a Mylar® collar, is used to apply radial compression to the stent and the two ends of the suture are passed through a Delrin guide before passing through the rollers of the test fixture.
• the stent samples were compressed and held at the predetermined diameter for approximately one minute, and then they were allowed to relax.
• the stents were subjected to two cycles of compression, hold and relaxation.
• the force exerted by the stent during the relaxation stage of the first cycle was recorded as the self-expansion force.
• the force applied to compress the stent in the second cycle was recorded as the compression resistance of the stent.
• the bioresorbabie stents of the present invention displayed substantially higher radial mechanical properties as compared to the Urolume® stent.
• Fig. 17 graphically depicts the cross-sectional lumenal area as a function of bilateral compression force for bioresorbabie fenestrated tube stents and 145° Urolume® stent.
• Fig. 17 shows that for the same amount of bilateral compression, the reduction in the lumen size of a Urolume® metallic stent was significantly greater than that of the bioresorbabie stent 50 of the present invention.
• FIGS. 18 and 19 are bar charts that illustrate the compression resistance and self-expansion force as a function of in vitro aging for four bioresorbabie fenestrated tube stents.
• the four test groups were subjected to different combinations of annealing and sterilization.
• Figs. 18 and 19 show that all four test groups maintained approximately 80% to 95% of initial compression resistance and 88% to 100% of self- expansion force after three weeks of aging.
• Figs. 18 and 19 show that the annealed stents had approximately 18% to 23% higher initial compression resistance and approximately 25% to 45% higher initial self-expansion force than non- annealed stents.
• Figs. 13 and 14 also show that ethylene oxide (eto) sterilization provides some slightly increased mechanical properties.
• a non-toxic radio-opaque marker is incorporated into the polymer blend prior to extruding the monofilaments used to weave the stent.
• suitable radio-opaque markers include, but are not limited to, barium sulfate and bismuth trioxide in a concentration of between approximately 5% to 30%.
• Table 1 represents the results obtained from testing different lots and configurations of the polymeric stents of the present invention. The stents polymers . and were tested as described in Examples 1 , 4 and 5.
• 1 2W, 3W, 4W, 5W and 6W refer to the number of week strength is retained.
• self-expanding stent may be utilized in the treatment of urethral stenoses. Accordingly, the present invention is not limited to that precisely as shown and described in the present invention.

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  • 2002-04-29 WO PCT/US2002/013686 patent/WO2002098476A1/en not_active Application Discontinuation
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