WO2014143866A1 - Fibres de type âme-gaine et leurs procédés de fabrication et d'utilisation - Google Patents

Fibres de type âme-gaine et leurs procédés de fabrication et d'utilisation Download PDF

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Publication number
WO2014143866A1
WO2014143866A1 PCT/US2014/028021 US2014028021W WO2014143866A1 WO 2014143866 A1 WO2014143866 A1 WO 2014143866A1 US 2014028021 W US2014028021 W US 2014028021W WO 2014143866 A1 WO2014143866 A1 WO 2014143866A1
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WIPO (PCT)
Prior art keywords
core
sheath
mesh
polymer
fibers
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PCT/US2014/028021
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English (en)
Inventor
Quynh Pham
Xuri Ray YAN
Abby DELEAULT
Toby Freyman
Joseph Lomakin
Gregory T. Zugates
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Arsenal Medical, Inc.
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Publication date
Application filed by Arsenal Medical, Inc. filed Critical Arsenal Medical, Inc.
Priority to JP2016502686A priority Critical patent/JP2016519222A/ja
Priority to CA2906074A priority patent/CA2906074A1/fr
Priority to EP14727275.1A priority patent/EP2971291A1/fr
Priority to AU2014228086A priority patent/AU2014228086A1/en
Publication of WO2014143866A1 publication Critical patent/WO2014143866A1/fr

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Classifications

    • D—TEXTILES; PAPER
    • D01—NATURAL OR MAN-MADE THREADS OR FIBRES; SPINNING
    • D01D—MECHANICAL METHODS OR APPARATUS IN THE MANUFACTURE OF ARTIFICIAL FILAMENTS, THREADS, FIBRES, BRISTLES OR RIBBONS
    • D01D5/00—Formation of filaments, threads, or the like
    • D01D5/0007—Electro-spinning
    • 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
    • A61L15/00—Chemical aspects of, or use of materials for, bandages, dressings or absorbent pads
    • A61L15/16—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons
    • A61L15/22—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons containing macromolecular materials
    • A61L15/24—Macromolecular compounds obtained by reactions only involving carbon-to-carbon unsaturated bonds; Derivatives thereof
    • 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
    • A61L15/00—Chemical aspects of, or use of materials for, bandages, dressings or absorbent pads
    • A61L15/16—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons
    • A61L15/22—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons containing macromolecular materials
    • A61L15/26—Macromolecular compounds obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds; Derivatives thereof
    • D—TEXTILES; PAPER
    • D01—NATURAL OR MAN-MADE THREADS OR FIBRES; SPINNING
    • D01F—CHEMICAL FEATURES IN THE MANUFACTURE OF ARTIFICIAL FILAMENTS, THREADS, FIBRES, BRISTLES OR RIBBONS; APPARATUS SPECIALLY ADAPTED FOR THE MANUFACTURE OF CARBON FILAMENTS
    • D01F8/00—Conjugated, i.e. bi- or multicomponent, artificial filaments or the like; Manufacture thereof

Definitions

  • the present disclosure relates, among other things, to core-sheath fibers, to methods of making core-sheath fibers and to devices and applications associated with core-sheath fibers.
  • Fibers and collections of fibers have been used as materials in various industrial applications, including applications in medicine and surgery ranging from sutures to wound dressings to skin grafts to arterial grafts, among many others. These applications are based on the unique properties of fibers as materials.
  • multicomponent fiber comprise (a) a polymeric core that comprises a core-forming polymer and (b) a polymeric sheath at least partially surrounding the polymeric core that comprises a sheath-forming polymer that is different than the core-forming polymer.
  • core-forming polymers include, for instance, crosslinked polysiloxanes and thermoplastic polymers, among others.
  • sheath-forming polymers include, for instance, solvent-soluble polymers, degradable polymers and hydrogel-forming polymers, among others.
  • the multicomponent fibers are formed using coaxial electrospinning techniques.
  • Still other aspects of the present invention pertain to meshes and other articles that are formed using the multicomponent fibers.
  • FIG. 1 shows a photomicrograph of a cross-section of the PLGA/PDMS sheath/core fibers formed in accordance with an embodiment of the invention.
  • FIG. 2 shows the PDMS fibers of FIG. 1 after sheath layer removal.
  • FIGS. 3A-3B show top-down and cross-sectional photomicrographs of PLGA/PDMS sheath/core fibers formed in accordance with an embodiment of the invention, both before sheath removal (FIGS. 3 A and 3C) and after sheath removal (FIGS. 3B and 3D).
  • FIG. 4 shows an image of water droplet (left) and an oil droplet (right), placed on a PDMS mesh in accordance with the present invention.
  • FIG. 5 is a stress-strain diagram illustrating mechanical properties of a PDMS mesh in accordance with the present invention as compared to a cast PDMS film.
  • FIGS. 6A-6B show cross-sectional photomicrographs of PLGA/PDMS sheath/core fibers that were electrospun at three differing sheath ore flow rates, in accordance with an embodiment of the invention.
  • FIG. 7 shows photomicrographs of PVP/PDMS sheath/core fibers formed in accordance with an embodiment of the invention, which show: (A) a cross-section of core-sheath fibers where the PVP cured at 100°C; (B) the same fibers as in (A) after they have undergone water extraction ; (c) a cross-section of core-sheath fibers where the PVP cured at 150°C; (D) the same fibers as in (C) after they have undergone water extraction.
  • FIG. 8 shows FTIR (Fourier transform infrared spectroscopy) scans of a pure PDMS film, a pure PVP film and a PVP PDMS sheath/core fiber formed in accordance with an embodiment of the invention (cured at 100°C), when dry and when wet.
  • FTIR Fastier transform infrared spectroscopy
  • FIG. 9 shows FTIR scans of a pure PDMS film, a pure PVP film and a PVP/PDMS sheath/core fiber formed in accordance with an embodiment of the invention (cured at 150°C), when dry and when wet.
  • FIG. 10 is a stress-strain diagram illustrating mechanical properties of PVP/PDMS sheath/core fibers formed in accordance with an embodiment of the invention (cured at 100°C and 150°C), when dry and when wet.
  • FIGS. 11 A and 1 IB shows balloon formed from a hydrated PVP-PDMS fiber mesh cured at 100°C, in accordance with an embodiment of the invention, at two levels of expansion.
  • FIG. 12 shows photomicrographs of fibers with a hydrophilic polyurethane (HLPU) sheath and a more hydrophobic polyurethane (HBPU) core, also referred to herein as
  • HLPU hydrophilic polyurethane
  • HBPU hydrophobic polyurethane
  • HLPU/HBPU sheath/core fibers formed at four HLPU:HBPU ratios, in accordance with an various embodiment of the invention.
  • FIG. 13 shows swelling and tensile strength as a function of HLPU content for meshes formed from HLPU/HBPU sheath/core fibers formed in accordance with various embodiments of the invention.
  • FIG. 14 shows swelling and shrinkage as a function of HLPU content for meshes formed from HLPU/HBPU sheath/core fibers formed in accordance with various embodiments of the invention.
  • FIG. 15 shows swelling for meshes formed from four different HLPU/HBPU sheath/core fibers formed in accordance with the invention (Formulations A-D), as well as two commercially available wound dressings.
  • FIG. 16 shows wet tensile strength for meshes formed from four different
  • HLPU/HBPU sheath/core fibers formed in accordance with the invention (Formulations A-D), as well as two commercially available wound dressings.
  • FIG. 17 shows shrinkage for meshes formed from four different HLPU/HBPU sheath/core fibers formed in accordance with the invention (Formulations A-D), as well as two commercially available wound dressings.
  • FIGS . 18 A and 18B show photomicrographs of a mesh formed from HLPU/HBPU sheath/core fibers before and after annealing, respectively, in accordance with an embodiment of the invention.
  • FIG. 19 shows phosphate buffered saline (PBS) retention for meshes formed from annealed (B Annealed) and non-annealed (B Normal) HLPU/HBPU sheath/core fibers formed in accordance with the invention, as well as two commercially available wound dressings.
  • PBS phosphate buffered saline
  • FIG. 20 shows shrinkage/expansion for meshes formed from annealed (B Annealed) and non-annealed (B Normal) HLPU/HBPU sheath core fibers formed in accordance with the invention, as well as two commercially available wound dressings.
  • FIG. 21 is a photomicrograph of HLPU/HBPU sheath/core fibers with encapsulated silver nanoparticles.
  • multicomponent fibers which comprise a polymeric core and a polymeric sheath at least partially surrounding (i.e., encapsulating) the core.
  • fibers are used synonymously to refer to elongated structures that differ only by size (with “microfibers” indicating fibers that have cross-sectional diameters on the order of microns to hundreds of microns, “nanofibers” indicating fibers that have cross-sectional diameters on the order of nanometers to hundreds of nanometers, and “fibers” indicating fibers of any size).
  • Fibers in accordance with the present disclosure can thus be formed in a wide variety of sizes.
  • Preferred overall fiber diameters range from 0.05 to 50 microns ( ⁇ ) (e.g., ranging from 0.05 to 0.1 to 0.25 to 0.5 to 1 to 2.5 to 5 to 10 to 25 to 50 microns), more preferably 0.1 to 20 microns, among other possible dimensions.
  • Preferred core diameters range from 0.01 to 10 microns (e.g., ranging from 0.01 to 0.025 to 0.05 to 0.1 to 0.25 to 0.5 to 1 to 2.5 to 5 to 10 microns), among other possible dimensions.
  • Preferred sheath thicknesses range from 0.02 to 25 microns (e.g., ranging from 0.02 to 0.05 to 0.1 to 0.25 to 0.5 to 1 to 2.5 to 5 to 10 to 25 microns), more preferably ranging from 0.2 to 18 microns, among other possible dimensions.
  • the ratio of the sheath volume to core volume can vary widely.
  • Preferred sheath volume:core volume ratios range, for example, from 100:1 to 1 :100, among other values, for example ranging from 100: 1 to 50:1 to 25: 1 to 10:1 to 5: 1 to 2:1 to 1 : 1 to 1:2 to 1 :5 to 1 :10 to 1 :25 to 1 :50 to 1 :100.
  • Multicomponent fibers in accordance with the present disclosure can be formed using various fiber spinning techniques, including various melt spinning and solvent spinning methods.
  • solvent spinning techniques and more particularly, electrostatic solvent spinning techniques, are detailed herein, the invention is not limited to such techniques.
  • Further exemplary techniques for forming multicomponent fibers include hot melt spinning, melt electrospinning, centrifugal fiber spinning, wet spinning, dry spinning, gel spinning, gravity spinning, extrusion, extrusion spinning, and rapid prototyping, among others.
  • multicomponent fibers may be formed that comprise (a) a polymeric core that comprises a core-forming polymer and (b) a polymeric sheath at least partially surrounding the polymeric core that comprises a sheath-forming polymer that is different than the core-forming polymer.
  • Electrospinning is a process that uses an electrical charge to draw very fine, typically micro- or nano-scale, fibers from a liquid. Solvent electrospinning utilizes an electrical force applied to a polymer solution to induce electrospinning jets. As streams associated with the jets travel in the air (or other atmosphere), evaporation of the solvent results in a single long polymer fibers deposited on a grounded collector. The collected fibers can result in the formation of a mesh which may be used in various technologies in medical and non-medical industries including, for example, drug delivery devices, tissue engineering, nano-scale sensors, wound dressings, self-healing coatings, and filters, among many others.
  • a "mesh” is a structure that is formed by a collection of one or more fibers interlaced to form a three dimensional network. Meshes include woven and non- woven meshes.
  • Meshes in accordance with the present disclosure can vary widely in thickness with preferred thicknesses ranging from 10 to 5000 microns (e.g., ranging from 10 to 25 to 50 to 100 to 250 to 500 to 1000 to 2500 to 5000 microns), among other values.
  • the meshes of the present disclosure can vary widely in porosity.
  • the meshes of the present disclosure have a porosity of 99% or less, for example, ranging from 99% to 90% to 80% to 70% to 60% to 50% to 40% to 30% to 20% to 10% or less.
  • Porosity can be measured by determining the volume of the polymer and dividing that quantity by the volume of the mesh.
  • Polymer volume Mesh mass ⁇ Polymer density
  • Mesh volume Mesh length x Mesh width
  • x Mesh thickness Mesh area x Mesh thickness
  • Mesh porosity (Mesh volume - Polymer volume) ⁇ Mesh volume.
  • the porosity of a given mesh may be reduced by annealing the mesh at a temperature and for a time wherein a decrease in mesh porosity is observed.
  • core-sheath electrospinning also referred to herein as coaxial electrospinning, uses two concentric needles to separately deliver two solutions, specifically, an inner core polymer solution and an outer sheath polymer solution.
  • the core solution is delivered through the inner needle whereas the sheath solution is delivered through the outer needle.
  • the two different polymer solutions are ejected in a continuous stream toward a grounded collector; this forms a single core-sheath Taylor cone at the needle tip, leading to the formation of a core-sheath fiber.
  • the creation of core-sheath fibers using needles has limited throughput.
  • core-sheath fibers are generated using a high-throughput core-sheath needleless electrospinning fixture, which utilizes one or more slits on the surface of a hollow vessel to co-localize numerous materials to multiple sites that form Taylor cones, thereby promoting the formation of multiple electrospinning jets and thus multiple electrospun fibers.
  • the slits on the surface of the hollow vessel thus may generate high-throughput production of core-sheath fibers.
  • each jet that forms thus leads to one long continuous fiber that gets collected.
  • the collected mesh is therefore comprised of approximately 10 very long fibers intertwined with one another.
  • hundreds of jets form and disappear with each rotation of the drum.
  • the resulting mesh consists of thousands of relatively short fibers.
  • the design of the needleless electrospinning fixture takes into account processing parameters that may enable greater control over fiber diameter. For example, in addition to the solution properties, solution flow rates can be manipulated to control fiber diameter.
  • the number of jets produced can also be controlled, which may lead to differences in fiber diameter.
  • the fibers of any embodiment of the present disclosure may thus be collected in a non-woven mesh form.
  • alternate embodiments include fibers that are collected as aligned fibers (as through gap alignment or rotating drum), twisted yarns, ropes, in a pattern, or any other method of fiber collection known in the art of electrospinning.
  • Various aspects of the invention pertain to multicomponent fibers that are formed using silicone polymers (also referred to herein as “silicones”, “siloxane polymers” or
  • multicomponent fibers are formed that comprise (a) a polymeric core that comprises one or more silicone polymers and (b) a polymeric sheath at least partially encapsulating the core that comprises one or more additional polymers other than silicone, or vice versa.
  • the present disclosure is applicable to all siloxanes (i.e., compounds with -Si-O-Si linkages), including polysiloxanes, which are formed from multiple siloxane units,
  • Ri and R 2 are organic radicals, for example, linear, branched or cyclic alkyl groups (e.g., methyl groups, ethyl groups, propyl groups, isopropyl groups, butyl groups, isobutyl groups, sec-butyl groups, tert-butyl groups, cyclohexyl groups and so forth), which may be substituted or unsubstituted, as well as substituted or unsubstituted aryl groups (e.g., phenyl groups, p-, m- or o-alkyl-substituted phenyl groups, and so forth).
  • Ri and R 2 can be the same or different.
  • polysiloxanes including as PDMS can be functionalized by a variety of mechanisms (e.g. plasma, UV, CVD, etc.) to modify the surface properties (e.g.
  • Fibers can be functionalized resulting in immobilized biomolecules on the surface and/or in the bulk. Functionalization can provide many new properties to the material, including biological effects, sensor applications. Microfibers and nanofibers further enhance these benefits by providing high surface areas and small pores, for example.
  • a functionalizing moiety for the PDMS is incorporated into the fiber. Upon curing, the functional moiety in the fiber becomes incorporated into the PDMS through siloxane chemistry. This allows for one-step functionalizing of the PDMS.
  • PDMS surfaces can be
  • polydimethylsiloxane for protein immobilization
  • PDMS polydimethylsiloxane
  • biotin groups can then be further modified with avidin-conjugated to a species of interest, for example, proteins, antibodies or fragments thereof, to functionalize the silicone surface. This may be useful, for example, in removing proteins from a liquid (e.g. protein separation) or in medical implants where preferential binding of certain proteins is advantageous (e.g. improved endothelial cell interactions).
  • a species of interest for example, proteins, antibodies or fragments thereof
  • This may be useful, for example, in removing proteins from a liquid (e.g. protein separation) or in medical implants where preferential binding of certain proteins is advantageous (e.g. improved endothelial cell interactions).
  • some classes of polymers including various siloxane polymers, are difficult to electrospin due to their low molecular weight and flowability. In this regard, various polysiloxanes remain flowable until they are crosslinked, which does not allow for sufficient poly
  • polydimethyl siloxane is a silicon-based organic polymer belonging to a larger group of siloxane polymers as indicated above, which commonly exhibit properties of elasticity and durability.
  • PDMS polydimethyl siloxane
  • the ability to manufacture fibers and constructs made from PDMS and other siloxane polymers that exhibit such properties, along with an ability to control the fiber diameter, is highly advantageous in medical technologies as well various other applications.
  • electrospin PDMS fibers the techniques developed thus far use blended polymer systems (i.e. not pure PDMS) and there are currently no electrospinning methods known to the inventors for manufacturing pure PDMS fiber constructs such as meshes.
  • core-sheath electrospinning techniques are provided, which can be used form fibers that comprise silicone materials that have not been previously electrospun using known techniques.
  • the fibers formed by the techniques described herein comprise a silicone material as the core material, and a different polymer material as the sheath material.
  • the core-sheath fibers are typically crosslinked by a suitable mechanism. For example, the fibers may be cured overnight at room temperature or for a few hours at temperatures up to 100°C, among other crosslinking techniques.
  • the polymeric sheath may be formed from hydrophilic or hydrogel materials, which are discussed in more detail below.
  • the polymeric sheath may be formed from materials that can be dissolved, degraded or otherwise removed from the silicone core, leaving behind pure silicone fibers.
  • materials include degradable polymers and solvent-soluble polymers, including water-soluble polymers.
  • Examples of degradable polymers include one or more of the following, among others: (a) polyester homopolymers and copolymers such as polyglycolide (PGA) (also referred to as polyglycolic acid), polylactide (PLA) (also referred to as polylactic acid) including poly-L- lactide, poly-D-lactide and poly-D,L-lactide, poly(lactide-co-glycolide) (PLGA), polycaprolactone, polyvalerolactone, poly(beta-hydroxybutyrate), polygluconate including poly- D-gluconate, poly-L-gluconate, poly-D,L-gluconate, poly(p-dioxanone), poly(lactide-co-delta- valerolactone), poly(lactide-co-epsilon-caprolactone), poly(lactide-co-beta-malic acid), poly(beta-hydroxybutyrate-co-beta-
  • water-soluble polymers include non-crosslinked hydrophilic polymers, which may be selected from homopolymers and copolymers formed from one or more of the following monomers, among others: ethylene oxide, vinyl pyrrolidone, vinyl alcohol, vinyl acetate, vinyl pyridine, methyl vinyl ether, acrylic acid and salts thereof, methacrylic acid and salts thereof, hydroxyethyl methacrylate, acrylamide, ⁇ , ⁇ -dimethyl acrylamide, N- hydroxymethyl acrylamide, alkyl oxazolines, saccharide monomers (e.g., polysaccharides such as dextran, alginate, etc.), and amino acids (e.g., hydrophilic polypeptides and proteins such as gelatin, etc.). When crosslinked, the preceding hydrophilic polymers are useful as hydrogels.
  • the preceding hydrophilic polymers are useful as hydrogels.
  • microarchitecture is highly dependent upon fiber diameter. Accordingly, an advantage of this core-sheath manufacturing process in which the sheath is subsequently removed is the ability to obtain pore sizes, porosities and other microarchitectural features.
  • the ratio of sheath-to-core thickness can be varied to provide larger pore sizes with smaller fibers or higher porosities with smaller fibers than can be obtained with other fabrication techniques. Fibers with Hydrogel Components and Components of varying
  • multicomponent fibers that are formed using hydrogels.
  • multicomponent fibers are formed that comprise (a) a polymeric core that comprises one or more core-forming polymers and (b) a polymeric sheath that comprises one or more hydrophilic or hydrogel-forming polymers.
  • Various aspects of the invention pertain to multicomponent fibers that comprise (a) a polymeric sheath that comprises one or more hydrophilic polymers and (b) a polymeric core that comprises one or more polymers that are more hydrophobic than the one or more hydrophilic polymers.
  • other aspects of the invention pertain to multicomponent fibers that comprise (a) a polymeric core that comprises one or more hydrophilic polymers and (b) a polymeric sheath that comprises one or more polymers that are more hydrophobic than the one or more hydrophilic polymers.
  • Polymers for use as core and/or sheath polymers include those that, upon immersion in an aqueous medium (e.g., water, PBS, etc.) at 25°C for one hour have water absorption values ranging anywhere from 0% to 1000% or more water, calculated as
  • a “hydrophilic polymer” is one that has a water absorption value ranging from from 5- 1000%» or more water.
  • a “more hydrophobic” polymer, also referred to herein as a “less hydrophilic” polymer, is defined as a polymer that absorbs less water than a given polymer to which it is being compared.
  • core and sheath polymers are selected such that the ratio of the sheath polymer water absorption value relative to the core polymer water absorption value ranges from 2:1 to 100:1 (for example ranging from 2:1 to 5:1 to 10:1 to 20:1 to 50:1 to 100:1), among other possible values, preferably 5: 1 to 20: 1 in certain embodiments.
  • the water absorption value of the sheath polymer in Example 4 below is 500% whereas the water absorption value of the sheath polymer is 50%, yielding a sheath ore water absorption ratio of 10:1.
  • Hydrogels comprise a three dimensional crosslinked network of hydrophilic polymers which have the ability to absorb substantial amounts of water. Hydrogels have long been used in in many applications in the medical field, ranging from drug delivery to tissue engineering scaffolds. Despite many potential applications, hydrogels have limited utility in healthcare or other fields due to a lack of structural control and a poor understanding of hydrogel mechanical properties. Others in the field have looked into reinforcing hydrogels with a variety of additives. Still others have aimed to reinforce hydrogels by making a polymeric fiber or polymeric fiber construct (e.g. a mesh) and then submersing it in a hydrogel or hydrogel-forming polymer before cross-linking the polymer. Such methods and structures have been generally ineffective, and there remains a need for hydrogel structures with desired properties.
  • a polymeric fiber or polymeric fiber construct e.g. a mesh
  • electrospinning is used to form a fiber core that comprises one or more fiber-forming polymers at least partially surrounded by a sheath that comprises one or more hydrogel-forming polymers.
  • the resulting composite fiber may be optionally subjected to a crosslinking step (e.g., by application of energy such as heat, visible light or ultraviolet light, by application of a crosslinking agent, etc.) to crosslink the hydrogel- forming polymers, the core-forming polymers, or both.
  • a crosslinking step e.g., by application of energy such as heat, visible light or ultraviolet light, by application of a crosslinking agent, etc.
  • the result is a composite fiber that has mechanical and hydration properties that differ from either material alone.
  • These composite fibers can be gathered, formed or processed into various shapes (e.g., tube, mesh, yarns, etc.) for use as medical devices or other products.
  • Polyurethanes may be employed as core and/or sheath polymers in various embodiments.
  • Polyurethanes are generally formed from diisocyanates and long-chain diols and, typically, chain extenders.
  • Aromatic diisocyanates may be selected from suitable members of the following, among others: methylenediphenyl diisocyanate (MDI) 5 toluene diisocyanate (TDI), naphthalene diisocyanate (NDI), para-phenylene diisocyanate (PPDI), 3,3'-tolidene-4,4'- diisocyanate and 3,3'-dimethyl-diphenylmethane-4,4'-diisocyanate.
  • Non-aromatic (aliphatic) diisocyanates may be selected from suitable members of the following, among others:
  • HDI hexamethylene diisocyanate
  • Hi 2 MDI dicyclohexylmethane diisocyanate
  • IPDI isophorone diisocyanate
  • CHDI cyclohexane diisocyanate
  • TMDI 2,2,4-trimethyl-l,6-hexamethylene diisocyanate
  • TMXDI meta-tetramethylxylyene diisocyanate
  • Long chain diols include polyether diols (e.g., polyethylene glycol, polyoxypropylene glycol, polytetramethylene ether glycol, etc.), polyester diols (e.g., polybutane diol adipate, polyethylene adipate, polycaprolactone diol, etc.), and polycarbonate diols.
  • Other long-chain diols include diols such as 1,4 butane diol, among others.
  • Polyurethanes other than those described in the prior paragraph may also be employed as core and/or sheath polymers in various embodiments
  • Hydrogels for use in the present disclosure include those formed from hydrophilic polymers which are crosslinked via a suitable mechanism, for example, covalently crosslinked and/or non-covalently crosslinked (e.g., by ionic crosslinking, physical crosslinking, etc.).
  • hydrophilic polymers which may be crosslinked include various hydrophilic polymers such as those set forth above.
  • hydrophilic polymers include hydrophilic polyurethanes (e.g., polyurethanes having hydrophilic segments), which may be physically crosslinked (e.g., via hard segments present in the polyurethanes).
  • hydrophilic polyurethanes include aliphatic, polyether-based polyurethanes and aromatic, polyether-based polyurethanes, among others. It is further noted that the hydrophilic polymers set forth above may be employed as hydrophilic segments in polyurethanes in certain embodiments.
  • core-forming polymers which include thermoplastic polymers and polymers of varying hydrophilicity/hydrophobicity in many embodiments, include silicones (polysiloxanes) such as those described above, thermoplastic polyurethanes such as aliphatic, polyether-based polyurethanes and aromatic, polyether-based polyurethanes, among others, and polyamides (e.g., nylon-6,6, nylon-6, nylon-6,9, nylon-6,10, nylon-6,12, nylon-11, nylon-12, nylon-4,6, etc.), among others.
  • core-forming polymers further include
  • homopolymers and copolymers comprising one or more of the following monomers, among others: (a) unsaturated hydrocarbon monomers (e.g., ethylene, propylene, isobutylene, 1-butene, 4-methyl pentene, 1-octene and other alpha-olefins, isoprene, butadiene, etc.); (b) halogenated unsaturated hydrocarbon monomers (e.g., tetrafluoroethylene, vinylidene chloride, vinylidene fluoride, chlorobutadiene, vinyl chloride, vinyl fluoride, etc.); (c) vinyl aromatic monomers including unsubstituted vinyl aromatic monomers (e.g., styrene, 2- vinyl naphthalene, etc.) and vinyl substituted aromatic monomers (e.g., alpha-methyl styrene), ring-substituted vinyl aromatic monomers; and (d) relatively hydrophobic (meth) unsaturated hydrocarbon mono
  • halo-alkyl (meth)acrylates e.g., 2,2,2-trifluoroethyl acrylate. It is noted that many of the preceding polymers can be employed as segments in polyurethanes in some embodiments.
  • Advantages associated with providing multi-component fibers with a hydrogel sheath and a core material that differs from the sheath material is that fibers, meshes and other constructions can be formed which have good water absorption and retention properties (as a result of the hydrogel material) coupled with desirable mechanical properties such as strength, elasticity, durability and shrinkage (as a result of the core material).
  • multicomponent fibers that comprise (a) a polymeric core that comprises one or more silicone polymers and (b) a polymeric sheath that comprises one or more additional polymers other than silicone.
  • the polymeric sheath may be formed from materials that can be dissolved, degraded or otherwise removed from the silicone core, leaving behind pure silicone fibers. Examples of such materials include degradable polymers and solvent-soluble polymers (including water soluble polymers) such as those set forth above, among others.
  • the fibers can be formed or processed into various shapes (e.g., tube, mesh, yarns) for use as medical devices or other products.
  • a silicon core-forming polymer is co-electrospun with a removable (e.g., dissolvable or degradable) sheath-forming polymer to create novel composite fibers.
  • the electrospinning may achieved by needleless electrospinning, coaxial electrospinning, slit-surface electrospinning, or any other suitable technique known in the art of fiber spinning.
  • fibers are formed with a PDMS core and a biodegradable polymer sheath.
  • Cross-linking of PDMS is performed using a two-part system by mixing the pre-polymer and a cross-linking agent which initiates the cross-linking reaction (exposure to heat accelerates this reaction).
  • a cross-linking agent which initiates the cross-linking reaction (exposure to heat accelerates this reaction).
  • two-part PDMS systems can be cured by exposure to UV-light.
  • two-part PDMS systems can be crosslinked into elastomers through free radical, condensation, or addition reactions.
  • one-part PDMS systems may be used which cure upon exposure to moisture in the atmosphere or photo- curing, among other techniques. Any of these variations in PDMS chemistries, or other polymers that require physical or chemical cross-linking to become a solid, may be used in the fibers and methods described herein.
  • a polysiloxane i.e., PDMS
  • PDMS polysiloxane
  • other embodiments may use polymers (e.g., thermosetting polymers, etc.) that require cross-linking to become solid.
  • polymers e.g., thermosetting polymers, etc.
  • examples include other polysiloxanes and certain types of polyesters, polyurethanes, polyimides, epoxies, etc.
  • degradable polymers i.e., poly(lactide-co-glycolides)
  • a solvent- soluble polymer sheath e.g., formed from a water-soluble sheath material such as uncrosslinked PEO, PVA, gelatin, dextran, carbohydrates, etc.
  • Embodiments employing aqueous solvents as dissolution agents generally do not result in swelling of PDMS fibers.
  • the sheath is etched away using an acid.
  • mechanical disruption may be used to break apart the sheath. Any combination of the described methods, or other suitable means, may be employed to remove the sheath from the underlying core.
  • therapeutic agents such as small molecule drugs, anesthetics, procoagulants, anticoagulants, antimicrobials, biologies, RNAi, genetic material, genetic vectors, vaccines, or particles such as silver nanoparticles are within the polysiloxane core.
  • a porogen e.g., selected from salts, sugars, etc.
  • porogen incorporated within the polysiloxane core. Upon subsequent sheath removal, the porogen is also removed. This will leave behind a fiber with porosity or rough surface features that may improve hydrophobicity, among other properties.
  • a porogen may be incorporated into the sheath such that after fiber formation, a certain percentage of the porogen is located at the interface of the core and sheath. Upon sheath removal, there is a negative imprint of the
  • the surface of polysiloxane fibers can also be roughened by a suitable etching process (e.g., laser etching) or mechanical means.
  • porosity can be introduced to the fibers of the present disclosure as a product of the cross-linking reaction that forms the fiber.
  • isocyanate functionalized PDMS can react with water to form porous foam fibers.
  • the acetic acid byproduct of the cross-linking reaction can react with sodium bicarbonate can generate gas and porosity, thus allowing for the formation of porous foam fibers.
  • Manipulation of fiber size can yield different fiber properties. For example, in filtration applications, smaller fibers with larger pores or higher porosity can increase the permeability and surface area.
  • Polysiloxane materials e.g., PDMS
  • PDMS polysiloxane materials
  • polysiloxane fiber meshes formed in accordance with the present disclosure have high surface area due to the small size fibers, which can promote adhesion and wetting where desired. In some embodiments, these same properties may be useful in medical applications where cell infiltration into fibers is desired.
  • smaller fiber diameters generally facilitate cellular interaction, ingrowth and proliferation while larger pores and higher permeability generally facilitate nutrient, cytokine and gas exchange while also improving cell migration.
  • the sheath is left on the core in order to form a composite fiber that contains a PDMS core and polymer sheath (e.g., nylon, polyethylene, polystyrene, polycarbonates, etc.) that possesses unique properties.
  • polymer sheath e.g., nylon, polyethylene, polystyrene, polycarbonates, etc.
  • the outer sheath may form a hydrogel to fill the porosity of a PDMS fiber mesh.
  • core and sheath polymers are reversed, and a polysiloxane is used as the sheath polymer that coats a core polymer. This allows the formation of bi- component fibers with the polysiloxane on the outside. Additionally, removal of the core polymer results in polysiloxane hollow fibers.
  • Small diameter fiber meshes can provide higher surface area, higher permeabilities and lower pore sizes than meshes made from larger diameter fibers.
  • the present disclosure thus provides materials which combine the benefits of polysiloxanes such as PDMS and small- diameter fiber meshes.
  • solvent-resistant filters or elastomeric, biocompatible for example, solvent-resistant filters or elastomeric, biocompatible
  • microfiber or nanofiber medical device components e.g., heart valve leaflets, vascular grafts, stent graft coverings
  • vascular grafts vascular grafts
  • stent graft coverings vascular grafts
  • silicone meshes may be used in heart valve leaflets.
  • Replacement heart valves in some cases, use synthetic materials to recreate the native leaflets.
  • Native leaflets are thin, highly flexible and durable. In addition to these properties the leaflets need to be nonthrombogenic. Encouraging endothelialization is one of the best ways to provide nonthrombogenic implants.
  • the microfiber architecture provided by electrospun silicone is thought to encourage endothelial cell growth. However, this same porosity may lead to blood passing through the pores of the mesh and reduced blood flow control by the valve. This phenomenon is expected to be temporary, however, as proteins and cells become trapped in the pores.
  • microfiber meshes of silicone are electrospun to a thickness of between 100 to 1000 microns (um).
  • Target fiber diameters are between 500 nm to 10 um. These meshes are then cut into appropriate shapes and attached to a main body which will be implanted via open or minimally-invasive surgery.
  • Alternate embodiments include: providing a membrane (e.g., silicone, PLGA) either on one side of the mesh or sandwiched between two meshes to prevent blood flow through the mesh; functionalizing the silicone with proteins or antibodies (e.g., CD34, VEGF) to encourage tissue ingrowth and reendothelialization; electrospinning onto a frame (e.g., polymer fiber, metal wire, contoured conductive mesh) to help shape the leaflet and/or provide an attachment to the main body; electrospinning onto a biocompatible fiber structure which will create a composite implant (e.g., fibers provide additional mechanical strength or varying stiffness across the leaflet); and coating or
  • a membrane e.g., silicone, PLGA
  • thrombogenicity e.g., heparin
  • silicone meshes may be used in stent graft coverings in a method similar to the heart valve leaflet, except that the silicone fibers are electrospun onto a tubular collector to form a tube of silicone microfibers or nanofibers.
  • Preferred mesh thickness is between 100 and 1000 microns.
  • Target fiber diameters are between 500 nm to 10 um.
  • This tube can then be attached to a stent to form the stent graft.
  • the fibers may be electrospun directly onto the stent.
  • Advantages include the fact that silicone microfibers and or nanofibers will encourage cellular ingrowth while providing an elastic, biocompatible, durable implant.
  • silicone meshes may be used in vascular grafts similar to the stent graft
  • the tube is not attached to a stent and the preferred mesh thickness range is larger (100 to 5000 microns).
  • silicone meshes may be used in bioengineered blood vessels.
  • the silicone mesh may be fashioned into a tube and seeded with cells ex vivo. These cells, typically fibroblasts, smooth muscle cells and endothelial cells, are incubated under various conditions (e.g., pulsatile flow, steady flow, no flow) in nutrient-rich environments to grow tissue on the graft material.
  • the silicone microfibers and nanofibers provide advantages in encouraging cell infiltration and growth as well as provide an elastic character typical of blood vessels.
  • the silicone tube may be used alone or in combination with other natural (e.g., collagen) or synthetic (e.g., PTFE, ePTFE, polyurethane) materials.
  • the graft is seeded with cells and implanted without significant incubation or implanted without cell seeding. In the latter case, cells from the host will infiltrate and populate the graft.
  • silicone meshes may be used in arteriovenous (AV) grafts and shunts. These grafts are used in hemodialysis patients to provide better needle access for repeated dialysis. Silicone microfiber or nanofiber meshes will provide a robust set of mechanical properties as well as encourage cellular ingrowth. The elasticity, durability, biocompatibility and low thrombogenicity of silicone will improve the performance of these grafts.
  • a silicone microfiber or nanofiber mesh is fashioned into a tube and implanted. This tube may be pre-treated by functionalization or coating with other materials (e.g., heparin, collagen, gelatin, growth factors) to improve integration and cell ingrowth.
  • the silicone mesh may be combined with other natural or synthetic materials as sheets or meshes to form a composite, layered structure. This layered structure may improve the mechanical properties, the ability to contain blood immediately after implantation or long term durability or performance.
  • silicone fibers electrospun into a flat mesh configuration of thickness 500 to 5000 microns may be used in hernia meshes.
  • a composite may be formed with biocompatible polymer fibers by electrospinning directly onto those fibers in the desired configuration. These fibers may also be provided in a configuration that improves suture-ability of the mesh.
  • the mesh may
  • silicone meshes may be used in dural covering.
  • a silicone microfiber or nanofiber mesh optionally combined with a polymer membrane (e.g., silicone, PLGA, collagen) can be used for this purpose.
  • silicone meshes may be used in wound dressing. Challenges for wound dressings include adherence to the wound and permeability to air and water (wound exudates).
  • silicone is electrospun into a mesh between 100 and 5000 microns thick. Preferred fiber diameters are between 500 nm and 10 microns. The electrospun silicone is non-adherent to the wound and provides high permeability and will be used a wound contacting layer in a dressing.
  • the silicone dressing is supplied separately and medical staff may place additional gauze or other bandages in layers on top of the silicone dressing.
  • the silicone mesh is combined with a gauze or other backing material as part of the finished product to absorb fluid and protect the wound.
  • the silicone can be fabricated with therapeutic agents such as antibiotics, antifungals, topical pain relievers, disinfectants (e.g., iodine) or the like.
  • Another embodiment provides a silicone mesh that has been treated with or manufactured with a hydrogel sheath (e.g., PEG) to provide moisture to the wound bed.
  • a hydrogel sheath e.g., PEG
  • the silicone mesh is fabricated for use with negative pressure wound therapy. In this case, the mesh is sized to be compatible with these devices and is placed on the wound bed as negative pressure is applied.
  • the silicone fibers may be electrospun onto a collector with a shape and topography similar to the intended treatment site (e.g., face, hand, etc.). In this way, the dressing can improve the therapy by improved conformance to the wounded tissue.
  • silicone meshes may be used in hemostatic applications.
  • the device is configured much like the wound dressings, but the silicone microfibers or nanofibers are fabricated or surface modified with a prothrombotic or
  • procoagulant agent e.g., thrombin, kaolin, chitosan, fibrin, etc.
  • the silicone provides a nonadherent dressing that can be removed easily.
  • the high permeability and porosity allows the blood to penetrate and contact a large amount of the surface area with the
  • prothrombotic agent This open structure also allows for coagulation factor diffusion back into the wound promoting clot formation.
  • This material may also be integrated as a non-adherent layer on other dressings (e.g., Combat Gauze); for this application the fibers may or may not be manufactured with a prothrombotic or procoagulant agent.
  • silicone meshes may be used in filtration applications.
  • Silicone meshes may be used as filters or as part of a filter for air, other gases, liquids, slurries or particles.
  • the high solvent resistance and durability provide advantages over other microfiber and nanofiber filters.
  • the low pore size and high permeability of electrospun, microfiber nanofiber meshes are desirable for filters.
  • the elastomeric nature provides a way to clean the filter. Simply stretching the material biaxially, circumferentially or otherwise will increase the pore size. Then, backflow of gas or liquid will provide a method to clear the pores of debris or other material. In a similar manner, cake which forms on the intake side of a filter may be easily removed by stretching the silicone mesh allowing the cake to fall off.
  • the silicone microfiber or nanofiber mesh may be used alone (preferred thickness of 100 microns to 1 cm).
  • the silicone mesh may be constructed as part of a layered filter using other commonly available filter materials.
  • the silicone may be electrospun directly onto another material, placed on the other material during assembly or electrospun onto a wire or other fiber mesh with large openings to provide mechanical support.
  • silicone meshes may be used in drug delivery.
  • Drugs may be incorporated into the silicone microfibers or nanofibers for delivery to a patient.
  • a silicone mesh is formed with drug in the silicone solution and is placed on the skin for cutaneous or transcutaneous delivery.
  • the silicone microfibers or nanofibers are formed into a mesh, tube or other structure and implanted to deliver drugs internally. This could include the mouth or other bodily orifices (e.g. , delivery of fluoride, bleach or other whitening substances to teeth).
  • silicone meshes may be used in barriers to modulate water penetration for controlled drug delivery.
  • a mesh of polysiloxane fibers such as silicone fibers could act as a barrier to modulate drug release. For example, if a drug delivery device has a
  • a PDMS mesh (which is relatively hydrophobic) can be placed around the device to prevent or slow water contact with the device. Additionally, since silicone is elastic, expansion of the mesh can lead to changes in its porosity and pore size, resulting in an increase of water so as to cause more drug release.
  • silicone meshes may be used in pressure-sensitive adhesive bandages.
  • the silicone microfibers or nanofibers are electrospun from a silicone which has adhesive properties. The mesh can then be applied to skin and will adhere well, but will provide water and air permeability to facilitate natural skin function and health.
  • This material can be used in bandages, as part of a wound dressing or for drug delivery patches.
  • silicone meshes may be used for oil-water separation as silicone is known to be relatively hydrophobic. With high pore volume fraction, a silicone microfiber or nanofiber mesh will separate oil from water.
  • the silicone may be surface treated, functionalized or doped with additives to make it more oleophilic or hydrophobic.
  • the silicone mesh may be used as a filter or placed into oil- water mixtures to remove oil or to separate oil from water. This may be extended to other systems containing hydrophilic and hydrophobic materials or phases. Because the mesh is highly elastic, the mesh can be stretched, squeezed, or compressed to clean/remove the oil from the pores for recovery of the oil and/or reuse of the mesh. Additionally, silicone also absorbs organic solvent and can also be used to separate aqueous from organic solvents. The high surface area of microfiber meshes makes it particularly efficient and appealing for these applications.
  • silicone meshes may be used in textiles. Silicone microfibers or nanofibers may also be used in textile applications where high elasticity, durability and permeability is desired. In other applications, the hydrophobicity or liquid repellent nature of silicone microfiber or nanofiber meshes (due to architecture) can be used to provide protection from liquids while still allowing air permeability to enable the skin to "breath”.
  • the composite fiber can be collected into aligned fiber bundles like a yarn.
  • These yarns will act as strong, elastic fibers that can be used (e.g., sutures) or processed further, including: twisting multiple yarns together into a rope, weaving multiple yarns together into a woven sheet, tube or other shape, braiding multiple yarns together into a stent, scaffold or other tubular structure.
  • Novel materials can be produced by forming various hydrophilic or hydrogel materials around various polymeric core materials, which act as a reinforcing material for the hydrophilic or hydrogel material.
  • the encapsulated polymer material can impart unique material properties (mechanical, chemical, thermal, etc.) to the hydrophilic or hydrogel material that would otherwise not be possible.
  • a core-forming polymer is co-electrospun with a hydrophilic or hydrogel-forming polymer to create novel composite fibers with a polymeric fiber core that is at least partially surrounded by a hydrophilic or hydrogel sheath.
  • the electrospinning may achieved by needleless electrospinning, coaxial electrospinning, slit-surface electrospinning, or any other suitable technique known in the spinning art.
  • the result is a composite fiber that has mechanical and hydration properties that are distinct from either material alone. These composite fibers can be gathered, formed or processed into various shapes (e.g., tube, mesh, yarns, etc.) for use as medical devices or other products.
  • any appropriate hydrophilic or hydrogel-forming material may be used as the sheath polymer and, like the selection of the polymeric core material, the hydrophilic or hydrogel- forming material can be selected to suit the particular purpose of the composite fiber.
  • the hydrophilic or hydrogel polymer sheath crosslinked PVP, PEO, PVA, and hydrophilic polyurethanes, among other polymers, as well as xerogels, aerogels, etc., may be used, among many other possibilities.
  • Other hydrogel polymers include crosslinked versions of hydrophilic polymers such as those listed above.
  • any appropriate polymer may be used for the core-forming polymer, depending on the mechanical or chemical needs at hand.
  • the fiber core is formed using a relatively hydrophobic polymer. While certain embodiments employ a covalently crosslinked silicon-based organic polymer core (e.g., a polysiloxane such as PDMS), the core polymer does not need to be covalently crosslinked to act as a reinforcing fiber.
  • thermoplastic polymers such as polyurethanes, PLGA, PCL, nylon, polystyrene, acrylic polymers, polypropylene, polyethylene and fluoropolymers, among others, can be used as the core reinforcing fiber.
  • Polyurethanes represent a broad class of polymers having a wide range of properties and, as such, can serve as core and/or sheath materials in conjunction with the present disclosure.
  • thermoplastic polyurethane core may be at least partially enclosed in a hydrophilic or hydrogel polyurethane sheath.
  • Many polyurethane materials exhibit physical cross-linking and thus do not require a separate crosslinking step. Such materials may be used, for example, in conjunction with melt-based or solvent-based spinning processes, among others.
  • polyurethane chemistry by co-electrospinning a hydrophilic polyurethane sheath around a more hydrophobic polyurethane core as detailed in Example 4 below.
  • the resulting composite fiber has mechanical and hydration properties that differ from either material alone.
  • a composite material consisting of a mechanically strong polyurethane core and a hydrophilic polyurethane sheath has been created.
  • the particular technique employed was slit-surface, core-sheath electrospinning.
  • electrospinning creates fibers with small diameters (micro or nanometers) which impart additional benefit and functionality (e.g., softness, high surface area, conformability).
  • suitable fibers may also be produced using other techniques including hot melt spinning, melt electrospinning, and centrifugal fiber spinning, among other fiber forming techniques.
  • the pre-polymer of PDMS is difficult to electrospin due to its low molecular weight and flowability, which does not allow for sufficient polymer chain
  • the silicone pre-polymer remains flowable until it is crosslinked, so spinning fibers without some way to preserve the fiber structure is unlikely to result in good fiber formation.
  • the present inventors have overcome this difficulty, particularly for micro and nano-sized fibers, by using coaxial electrospinning to encapsulate PDMS pre- polymer and a cross-linking agent within a polymer sheath.
  • a hydrogel polymer is used as a polymer sheath material.
  • the core polymer is crosslinked PDMS and the polymer sheath is a crosslinked polyvinylpyrrolidone (PVP).
  • the cross-linking of the hydrogel-forming polymer is modified to suit the core materials, as well as the desired properties of the composite fiber.
  • hydrogel crosslinking is initiated by the application of heat, along with core crosslinking.
  • the core polymer may be crosslinked PDMS and the polymer sheath may be a crosslinked polyvinylpyrrolidone (PVP), both of which are crosslinked by the application of heat (see, e.g., Example 3).
  • PVP polyvinylpyrrolidone
  • linking of the hydrogel polymer (and/or core polymer) could include UV or gamma radiation, freeze/thaw cycles, supercritical drying, and so forth.
  • a physically crosslinked hydrogel is selected (see, e.g., Example 4). All of these variations of hydrogel chemistries are within the present disclosure.
  • a major benefit of this aspect of the present disclosure is that an elastic, durable, biocompatible and mechanically stable construct may be provided for hydrogels so that the many potential benefits of hydrogels can be utilized in applications which require greater mechanical integrity.
  • Another benefit is that methods of forming core-hydrogel fibers are provided, which do not require a separate crosslinking step, due to the physical crosslinking attributes of the polymers selected as the core-forming polymer and/or sheath-forming polymer.
  • small diameter fiber meshes provide, inter alia, higher surface area, higher permeabilities and lower pore sizes than meshes made from larger diameter fibers. This disclosure thus provides materials which combine the benefits of hydrogels and small- diameter fiber meshes.
  • these core-hydrogel fibers can be gathered, formed or processed into various shapes (e.g., tube, mesh) for use as medical devices or other products.
  • excipient materials are incorporated into the fibers to increase water swelling and retention capacities.
  • Excipient materials include cross-linked hydrophilic polymers such as PVP, cellulose, gelatin and starch, among others. These materials can be incorporated as dissolved polymers in the sheath or core during electro spinning.
  • the excipient materials will present as particles embedded in or projecting from the surface of the finished fibers.
  • therapeutic agents such as small molecule drugs, anesthetics, procoagulants, anticoagulants, antimicrobials, biologies, RNAi, genetic material, genetic vectors, vaccines, or particles such as silver nanoparticles are incorporated into the fibers which are released upon hydration.
  • the composite core-hydrogel fiber can be used in heart valve leaflets.
  • Replacement heart valves in some cases, use synthetic materials to recreate the native leaflets.
  • Native leaflets are thin, highly flexible and durable. In addition to these properties the leaflets need to be non-thrombogenic. Encouraging
  • endothelial ization is one of the best ways to provide non-thrombogenic implants.
  • the hydrogel layer sheath along with the microfiber or nanofiber architecture will encourage endothelial cell growth. Upon hydration, the hydrogel layer will swell and fill the pores between the core fibers- -thus preventing blood from passing through the pores of the valve.
  • microfiber meshes of core-hydrogel fibers are electrospun to a thickness of between 100 to 1000 microns. Target fiber diameters are between 500 nm to 10 um. These meshes are then cut into appropriate shapes and attached to a main body which will be implanted via open or minimally- invasive surgery.
  • Alternate embodiments include: functionalizing the core polymer with proteins or antibodies (e.g. CD34, VEGF) to encourage tissue ingrowth and reendothelialization
  • electrospinning onto a frame e.g.
  • polymer fiber, metal wire, contoured conductive mesh to help shape the leaflet and/or provide an attachment to the main body; electrospinning onto a biocompatible fiber structure which will create a composite implant (e.g. fibers provide additional mechanical strength or varying stiffness across the leaflet); and coating or functionalizing the fibers to decrease thrombogenicity (e.g. heparin).
  • the composite core-hydrogel fiber can be used in stent graft coverings.
  • hydrogel fibers can be used as coverings on stents that are used in left atrial appendage closures. These embodiments are similar to the heart valve leaflet, but the core- hydrogel fibers are electrospun onto a tubular collector to form a tube of microfibers or nanofibers. Preferred mesh thickness is between 100 and 1000 microns. Target fiber diameters are between 500 nm to 10 um. This tube can then be attached to a stent to form the stent graft. Alternatively, the fibers may be electrospun directly onto the stent. Alternate embodiments described for the heart valve concept are applicable here as well. Advantages are that composite core-hydrogel fibers will encourage cellular ingrowth while providing an elastic, biocompatible, durable implant.
  • the composite core-hydrogel fiber can be used in vascular grafts. These embodiments are similar to the stent graft design but the tube is not attached to a
  • these tubular meshes act as a reinforcing cuff for vessels (e.g., vascular autografts for bypass surgeries) or other tubular structures where the mechanical properties of the native tissue have deteriorated, such as in abdominal aortic aneurysms.
  • vessels e.g., vascular autografts for bypass surgeries
  • other tubular structures where the mechanical properties of the native tissue have deteriorated, such as in abdominal aortic aneurysms.
  • the composite core-hydrogel fiber can be used in any embodiments.
  • bioengineered blood vessels are similar to the vascular graft above, and the core-hydrogel microfiber or nanofiber mesh can be fashioned into a tube and seeded with cells ex vivo. These cells, typically fibroblasts, smooth muscle cells and endothelial cells, are incubated under various conditions (e.g. pulsatile flow, steady flow, no flow) in nutrient-rich environments to grow tissue on the graft material.
  • the core-hydrogel microfibers or nanofibers may provide advantages in encouraging cell infiltration and growth as well as provide an elastic character typical of blood vessels.
  • the core-hydrogel tube may be used alone or in combination with other natural (e.g. collagen) or synthetic (e.g.
  • the graft is seeded with cells and implanted without significant incubation or implanted without cell seeding. In the latter case, cells from the host will infiltrate and populate the graft.
  • the hydrogel fibers are used in medical device sealing applications. These mechanically robust, hydrogel fibers and resulting meshes, yarns, tubes, etc. are ideally suited for use to seal interfaces between medical devices and the body, other medical devices or other surfaces requiring a seal. For example, they can be used to provide a seal between an implanted heart valve and the native valve annulus to prevent paravalvular leakage.
  • the hydrogel fibers are electrospun directly onto the outer surface of the valve stent or fashioned into a mesh, yarn or tube and applied to the valve stent as part of the manufacturing process.
  • hydrogel Upon implantation the hydrogel absorbs water from the blood which leads to swelling, filing of the space between the implant and the valve annulus and thus sealing around the valve to prevent leakage.
  • the advantage compared to other hydrogels is the favorable mechanical properties and durability lead to a safer and more effective product.
  • Other applications include: providing hydrogel microfibers or nanofibers on the vessel contacting side of a stent graft, vascular graft or other medical device to seal between the graft or other medical device and the vessel wall; providing hydrogel microfibers or nanofibers on the outer or inner diameter of a stent graft to seal between two stent graft components which will be assembled
  • EVAR graft main body and iliac limb extension e.g., EVAR graft main body and iliac limb extension
  • providing fibers on the outside of a stent graft to be used as a chimney, snorkel, etc. as part of another stent graft placement providing hydrogel microfibers or nanofibers on the outer surface of a transcutaneous catheter, ostomy bag, or wire lead to seal between the device and the skin and/or underlying muscle, fat or fascia
  • the robust mechanical properties and slippery surface of the hydrogel will aid in removal.
  • the hydrogel fibers can be manufactured such that they hydrate only when a strain is applied (see, e.g., Example 3 below). Upon hydration, the fibrous construct increases in volume. This property can be applied to create strain-dependent seals around stent grafts and heart valve cuffs. In some cases, when stent grafts and heart valve cuffs are deployed, they do not make complete conformal contact with the vessel wall or annulus, thereby leaving open spaces between the stent graft and vessel, which in turn may lead to leaks, device failure and poor clinical outcomes.
  • the hydrogel fibers can be used as a ring or stent covering such that during delivery, the hydrogel fiber covering does not wet, but upon stent deployment the fiber covering is strained, resulting in wetting and swelling of the fibers that fill empty spaces where the stent does not make conformal contact with surrounding tissues.
  • the hydrogel fibers are used in non-medical sealing.
  • the core-hydrogel fibers will be useful in providing a seal in non-medical applications in aqueous or non-aqueous environments.
  • fibers positioned between two surfaces to be sealed will hydrate upon contact with water then the swelling will seal the surfaces and prevent flow through the microstructure.
  • non-aqueous applications e.g., oil transport
  • the mesh will be hydrated upon installation creating a seal from swelling in between two surfaces and also prevent leakage due to immiscibility with the nonaqueous fluid.
  • the composite core-hydrogel fiber can be used in any embodiments.
  • arteriovenous grafts or shunts arteriovenous grafts or shunts. These grafts are used in hemodialysis patients to provide better needle access for repeated dialysis.
  • a core-hydrogel microfiber or nanofiber mesh will provide a
  • a core-hydrogel microfiber or nanofiber mesh is fashioned into a tube and implanted. This tube can be pre-treated by functionalization or coating with other materials (e.g. heparin, collagen, gelatin, growth factors) to improve integration and cell ingrowth.
  • a core-hydrogel mesh can be combined with other natural or synthetic materials as sheets or meshes to form a composite, layered structure. This layered structure may improve the mechanical properties, the ability to contain blood immediately after implantation or long term durability or performance.
  • the composite core-hydrogel fiber can be used in hernia meshes.
  • Core-hydrogel fibers e.g. , silicone or polyurethane core with a hydrogel sheath
  • a composite may be formed with biocompatible polymer fibers by electrospinning directly onto those fibers in the desired configuration. These fibers may also be provided in a configuration that improves suture-ability of the mesh.
  • the mesh may be functionalized to improve tissue ingrowth or integration.
  • the composite core-hydrogel fibers can be used in dural coverings.
  • a core-hydrogel microfiber or nanofiber mesh optionally combined with a polymer membrane (e.g. silicone, PLGA, collagen) can be used for this purpose.
  • a polymer membrane e.g. silicone, PLGA, collagen
  • the composite core-hydrogel fibers can be used in wound dressing.
  • Challenges for wound dressings include adherence to the wound, wound exudate management and permeability to air and water (wound exudates).
  • hydrogel encapsulated polymer e.g., silicone or polyurethane
  • Preferred fiber diameters are between 500 nm and 10 microns.
  • the advantage of the reinforced hydrogel is that it provides moisture to the wound bed while also forming a protective layer which does not adhere to the wound.
  • a hydrogel- polymer dressing is supplied separately and medical staff place additional gauze or other bandages in layers on top of the core-hydrogel fiber dressing.
  • a core- hydrogel mesh is combined with a gauze or other backing material as part of the finished product
  • a core- hydrogel mesh can be fabricated with therapeutic agents such as antibiotics, antifungals, topical pain relievers, disinfectants (e.g. iodine) or the like.
  • a hydrogel- polymer mesh is fabricated for use with negative pressure wound therapy. In this case, the mesh is sized to be compatible with these devices and is placed on the wound bed as negative pressure is applied. The high permeability and porosity allow exudates removal as well as a non-adherent dressing when it must be removed.
  • the hydrogel sheath or core polymer may also be useful in controlling release of therapeutic agents to the wound (e.g., antimicrobials, antibiotics, silver ions, growth factors, analgesics, anesthetics, debridement compounds or enzymes, etc.).
  • therapeutic agents e.g., antimicrobials, antibiotics, silver ions, growth factors, analgesics, anesthetics, debridement compounds or enzymes, etc.
  • the composite core-hydrogel fiber can be used in hemostat applications.
  • the device is configured much like the wound dressings, but the hydrogel- polymer microfibers or nanofibers are fabricated or surface modified with a prothrombotic agent (e.g. thrombin, kaolin, chitosan, fibrin).
  • a prothrombotic agent e.g. thrombin, kaolin, chitosan, fibrin.
  • the fiber provides a nonadherent dressing that can be removed easily.
  • the high permeability and porosity allows the blood to penetrate and contact a large amount of the surface area with the
  • prothrombotic agent This open structure also allows for coagulation factor diffusion back into the wound promoting clot formation.
  • This material may also be integrated as a non-adherent layer on other dressings (e.g. combat Gauze);for this application the fibers may or may not be manufactured with a prothrombotic agent.
  • the composite core-hydrogel fiber can be used in filtration.
  • Composite core-hydrogel fiber meshes can be used as filters or as part of a filter for air, gases, liquids, slurries or particles.
  • the low pore size and high permeability of electrospun, microfiber or nanofiber meshes are desirable for filters.
  • the fibers are elastomeric, the elastomeric nature provides a way to clean the filter. Simply, stretching the material biaxially, circumferentially or otherwise will increase the pore size. Then, backflow of gas or liquid will provide a method to clear the pores of debris or other material.
  • the core-hydrogel microfiber or nanofiber mesh may be characterized by high strength and hydrophilicity, thus being useful as a filter, barrier or separating membrane to partition oil content in water.
  • the core-hydrogel microfiber or nanofiber mesh can be used alone (preferred thickness of 100 microns to 1 cm). Alternately, the core-
  • the 29 hydrogel mesh can be constructed as part of a layered filter using other commonly available filter materials.
  • the core-hydrogel may be electrospun directly onto another material, placed on the other material during assembly or electrospun onto a wire or other fiber mesh with large openings to provide mechanical support.
  • the composite core-hydrogel fiber can be used in drug delivery.
  • the hydrated core-hydrogel composite material may act as a substantially non-porous yet conformal layer.
  • the core-hydrogel material would be inserted into the target delivery area then inflated with gas or other fluid (e.g., a drug containing solution, etc.) to conform to the internal structure of the target area. Direct, conformal contact of the hydrogel with the surface leads to efficient drug delivery.
  • gas or other fluid e.g., a drug containing solution, etc.
  • the pores upon reaching a certain expansion limit on inflation, the pores become stretched and open to allow drug solution to be released. Once deflated, the pores seal back up thus inhibiting drug delivery to areas not being targeted during removal of the device. This approach is particularly applicable for therapeutic delivery to cavities and lumens, such as the sinusoidal space.
  • drugs may be incorporated into the core-hydrogel microfibers or nanofibers for delivery to a patient.
  • a core-hydrogel fiber mesh may be formed with drug in the core-forming solution, and placed on the skin for cutaneous or transcutaneous delivery.
  • Fiber meshes of the present disclosure are beneficial in that they provide a means of targeted delivery to difficult orifices such as sinus cavities, intestinal wall or ear canals due to the ability to balloon open for conformal delivery.
  • a tubular or other shaped mesh may also be implanted to provide sustained drug delivery. It may be implanted alone or held in place using another medical device, such as a stent.
  • the composite core-hydrogel fibers can be collected into aligned fiber bundles like a yarn. These yarns will act as strong, elastic hydrogel fibers that can be used (e.g., sutures) or processed further, including: twisting multiple yarns together into a rope, weaving multiple yarns together into a woven sheet, tube or other shape, braiding multiple yarns together into a stent, scaffold or other tubular structure. These configurations can be developed into novel medical devices such as hydrogel catheters, introducer sheaths, guide wires, vascular grafts, hernia meshes, etc.
  • the composite core-hydrogel fiber can be used in textiles.
  • Core-hydrogel microfibers or nanofibers may also be used in textile applications where high elasticity, durability, water absorption and permeability are desired.
  • the composite core-hydrogel fiber can be used in tissue engineering applications. Hydrogels allow for free diffusion of oxygen, nutrients, etc., which is desirable for these purposes. This property is further enhanced, because diffusion not only can occur across the hydrogel bulk, but through the porosity created by the fibrous network.
  • Hydrogels are used extensively in tissue engineering applications due to their promising biocompatibility and hydration properties.
  • a major benefit of the present disclosure is that fibrous hydrogels would allow for better cell attachment and integration to form 3D scaffolds.
  • the hydrogel sheath would allow for cell attachment and in-growth, which could eventually degrade away, while the core polymer fibers would provide more permanent mechanical support.
  • a specific example application of this includes hyaline cartilage repair, in which the hydrogel sheath provides a biocompatible scaffold for stem cells to attach and differentiate into
  • chondrocytes while the porosity provides space for chondrocytic secretion of collagen and ECM components.
  • the hydrogel fibers are used as a tissue bulking agent in cosmetic or plastic surgery.
  • the elastic and flexible mechanical properties and high hydration of the hydrogel fibers can be tailored to match that of native tissue for a more natural look and feel.
  • the fibrous nature will integrate with the surrounding tissue such that the bulking agent stays in place and will not become displaced.
  • the hydrogel can be made to be
  • the composite core-hydrogel fiber are used as medical electrodes.
  • the swelling properties of hydrogel allow for conformable and intimate contact with tissue that can lower electrical impedance and improve electrode performance.
  • the core material can be comprised of a conductive polymer or include electrically conductive particles or ions.
  • the ballooning and hydration capability of the composite core-hydrogel fibers is a unique property that can be used for the ablation of tissues through the use of microwaves.
  • a mesh of composite core-hydrogel fibers can be used for ablation within a body cavity (e.g., endometrial, left atrial appendage) or to an irregular surface (e.g. , liver, esophagus, sinuses).
  • a gas e.g., carbon dioxide
  • microwaves from a source within the balloon will heat the water within the hydrogel membrane, which is in intimate and conformal contact with the cavity or tissue surface, to thermally ablate the surrounding tissue.
  • This same technique may be extended to other ablation approaches, including hydrothennal (e.g., inflate the balloon with hot water or other hot liquid), chemical (e.g., ablative agent in the hydrogel fibers) or cyroablation (e.g., cold source or liquid nitrogen used to chill the balloon).
  • hydrothennal e.g., inflate the balloon with hot water or other hot liquid
  • chemical e.g., ablative agent in the hydrogel fibers
  • cyroablation e.g., cold source or liquid nitrogen used to chill the balloon.
  • the composite core-hydrogel fibers of the present disclosure may also be used to embolize a body lumen.
  • the composite structure provides a fiber or coil that can be inserted into a patient using techniques know to those skilled in the art.
  • the hydrogel properties then swell the fibers to completely fill the body lumen or aneurysm cavity.
  • Two key advantages here are 1) combination of fiber strength and high swelling ratio, and 2) ability to form very small fibers or coils and / or flexible implants.
  • Example 1 Fibers with PDMS core and PLGA sheath.
  • Core/sheath fibers are fabricated in accordance using a high-throughput core-sheath needleless electrospinning fixture.
  • the sheath polymer system was a 3.5 wt% 85/15 poly(L- lactic acid-co-glycolic acid) (PLGA) in 6:1 (by vol) chloroform:methanol solvent.
  • the core polymer consisted of PDMS (Sylgard 184, available from Dow Corning, a two-part liquid system consisting of part A (pre-polymer) and part B (cross-linking agent)) mixed in a 10:1 mass ratio.
  • the sheath solution flow rate was set to 200 ml/h while the core flow rate was set to 20 ml/h.
  • FIG. 1 shows an image of the cross-section of the PLGA/PDMS sheath/core fibers after curing. The different polymers in the sheath/core configuration can be observed.
  • FIG. 2 shows the PDMS fibers after sheath layer removal. PDMS fibers were manufactured to be between about 1 and 5 microns in diameter. As described elsewhere herein, however, the diameter of the core PDMS can be tuned by modulating electrospinning parameters.
  • Core-sheath fibers were electrospun with 50/50 poly(D,L-lactic acid-co-glycolic acid) (5050 PDLGA) as the sheath over a PDMS (Sylgard 184) core, as described in Example 1.
  • the sheath solution was an 11 wt% 5050 PDLGA in hexafmoroisopropanol (HFIP).
  • the flow rate for the sheath solution was set at 10 ml/h while the core solution flow rate was set at 1 ml/h.
  • the fibers were subsequently placed in a 60°C oven for 24 hours to allow the PDMS in the fibers to cure.
  • FIG. 3A shows a core-sheath structure (in cross-section) that was formed.
  • Diameters of the fibers were measured for both top-down and cross-sectional images.
  • the overall fiber diameter of the fibers was approximately 7 microns (see FIGS. 3 A and 3C), while the core PDMS diameter was approximately 4.5 micron (see FIGS. 3B and 3D).
  • the 5050 PDLGA sheath was removed under accelerated degradation conditions by immersing the mesh in 12 pH buffer consisting of 1.5% sodium phosphate, 0.1% boric acid, and 0.08% citric acid at 37°C for 7 days. As can be seen in FIG. 3B, the sheath layer was completely degraded and removed, leaving behind PDMS-only fibers.
  • the electrospun fibers and meshes of the present disclosure offer different properties than those formed from traditional methods of constructing PDMS as a cast film.
  • the contact angle of the electrospun PDMS-only mesh was measured to be 110° while a cast film of PDMS had a contact angle of 104°.
  • FIG. 4 shows the hydrophobic and oleophilic nature of the PDMS mesh formed using the electrospinning processes of the present disclosure. A water droplet (left) placed on the mesh remains beaded while an oil droplet (right) wets the mesh and can move throughout the porosity of the mesh.
  • FIG. 5 shows the mechanical properties of the mesh compared to a cast PDMS film.
  • the modulus of the mesh is significantly lower (0.2 MPa vs 2.0 MPa) while its extension at max loading is significantly higher (300% vs. 122%) relative to the cast PDMS film.
  • FIGS. 6A-6C show cross-sectional photomicrographs of electrospun fibers of the present disclosure having PDMS in the core and 5050 PDLGA in the sheath.
  • electrospinning process was carried out at sheath.xore flow rates of 10:1, 10:0.25, and 20:0.25
  • Example 3 Fibers with PDMS core and PVP sheath.
  • Core/sheath fibers were fabricated using a sheath polymer solution of 8wt% PVP (polyvinyl pyrrolidone) in TFE (trifluoroethanol), while the core polymer solution consisted of Sylgard 184, a two-part liquid system consisting of Part A (pre-polymer) and B (cross-linking agent) mixed in a 10: 1 mass ratio.
  • the sheath flow rate was set to lOmL/h while the core flow rate was set to 2mL/h.
  • FIG. 7 shows: (A) SEM cross-section of core-sheath fibers where the core consists of fully cured PDMS and the sheath is PVP cured at 100°C; (B) SEM cross-section of the same fibers in (A) except after they have undergone water extraction to remove non-cross linked PVP; (C) SEM cross-section of core-sheath fibers where the core consists of fully cured PDMS and the sheath is PVP cured at 150°C; (D) SEM cross-section of the same fibers in (C) except after they have undergone water extraction to remove non-cross linked PVP. Before hydration, the two cure temperature samples look identical in core fiber diameter (around 6um) and in sheath
  • FIGS. 8 and 9 show the spectra obtained from the silicone fiber hydrogels in the wet and dry states as compared to pure PDMS and pure PVP cured at temperatures of 100°C and 150°C, respectively.
  • silicone core/hydrogel sheath fibers Another feature of silicone core/hydrogel sheath fibers is that its mechanical features can change, depending on whether the fibers are wet or dry. For example, a dry mesh will have high air permeability, high porosity and will be opaque. Conversely a hydrated mesh will have lower air permeability (because the swollen hydrogel fills the pores), high water permeability and will be optically clear. Upon reaching its expansion limit of the mesh the pores open up and the gas or liquid flows through the mesh as opposed to bursting it. This ability to expand is also affected by the cure temperature, because as the elongation of the fibers is dependent on cure temperature.
  • a "balloon" was formed from a mesh of hydrated PVP-PDMS fibers cured at 100°C. If spherical expansion is assumed, then the volume expansion ratio is near 800% when the balloon reaches maximum expansion (see Fig. 1 IB). At the maximum expansion, the balloon doesn't burst but merely becomes air permeable and allows air to escape through the pores. The air permeability and porosity of the hydrated mesh can be increased upon stretching the mesh to open up the pores. Due to the decreased permeability of the hydrated hydrogel mesh, the material can hold air or water and expand to very high volumes while still maintaining mechanical integrity. In addition, the microfibers provide a flexible balloon that can conform to irregular surfaces, cavities or containers.
  • the 100°C cured mesh absorbs nearly the same amount of oil as it does water due to the presence of the PDMS fibers and the very small amount of PVP on the surface.
  • the mesh cured at 150°C absorbs much more water than it does oil, because much more crosslinked PVP is present in this sample.
  • Example 4 Fibers with a poiyurethane core and a hydrophilic poiyurethane sheath.
  • slit-surface, core-sheath electrospinning was employed, in which a hydrophilic aliphatic polyether-based thermoplastic poiyurethane (HLPU) was used as the sheath material, while a mechanically stronger more hydrophobic aliphatic polyether-based thermoplastic poiyurethane (HLPU) was used as the sheath material, while a mechanically stronger more hydrophobic aliphatic polyether-based
  • thermoplastic poiyurethane material (HBPU) was used as the core material.
  • the electrospinning solutions were as follows: 4 wt% HLPU in TFE and 6 wt% HBPU in HFIP. Electrospinning was carried out at different sheath ore flow rate ratios. At the flow rate ratios selected, the resulting fiber was composed of HLPU and HBPU in the following HLPU:HBPU weight ratios: (A) 93:7, (B) 82: 18, (C) 60:40, and (D) 38:62, respectively.
  • Figure 12 shows the SEM of the fibers for each composition; fiber diameters for all formulations were approximately 2 microns
  • the wet tensile strength of the different polyurethane samples are shown in Table 2 and demonstrates an increase in mechanical properties as the amount of HBPU in the fiber is increased. Therefore, by varying the core to sheath material composition, one can modulate the tensile strength.
  • Table 3 shows the hydration properties of the different formulations and indicates that sample shrinkage upon hydration and swelling were most impacted by the chemical composition of the fibers. However, PBS retention did not appear to be significantly impacted.
  • FIGS. 13 and 14 A comparison of the mechanical and hydration properties as a function of HLPU content is shown in FIGS. 13 and 14.
  • FIG. 13 shows that tensile strength increases as the amount of HLPU decreases (and hence the amount of HBPU increases). However, as the tensile strength increases, the amount of PBS absorption decreases as a result of less hydrophilic material being present.
  • a comparison of the swelling (or PBS absorption) and shrinkage data as a function of the HLPU content further reinforces the utility of using a core-sheath fiber structure to modulate the mechanical and hydration properties.
  • FIG. 14 there is an increase in swelling capacity as the HLPU content increases; however, dimensional shrinkage (i.e., shrinkage in area) of the mesh is also observed to increase as the HLPU content increases.
  • meshes in accordance with the present disclosure are annealed at elevated temperature to improve the properties of the same. For example,
  • FIGS. 18A and 18B are photomicrographs of a mesh formed from HLPU/HBPU sheath/core fibers as described herein, before and after annealing, respectively.
  • the annealing step is accompanied by a reduction in mesh volume (and thus mesh area).
  • such an annealing step has been found to improve water retention and to result in mesh expansion (rather than mesh shrinkage).
  • FIG. 18A is a photomicrographs of a mesh formed from HLPU/HBPU sheath/core fibers as described herein, before and after annealing, respectively.
  • the annealing step is accompanied by a reduction in mesh volume (and thus mesh area).
  • such an annealing step has been found to improve water retention and to result in mesh expansion (rather than mesh shrinkage).
  • FIG. 19 shows PBS retention values for non-annealed (B Normal) and annealed (B Annealed) HLPU/HBPU sheath core fiber meshes in accordance with the present disclosure, as well as retention values for Aquacel® and Durafiber® wound dressings.
  • an annealed mesh material has been developed which provides PBS retention equivalent to that of Aquacel® and Durafiber® dressings.
  • FIG. 20 shows shrinkage or expansion values for non-annealed (B Normal) and annealed (B Annealed) HLPU HBPU sheath/core fiber meshes in accordance with the present disclosure, as well as for Aquacel® and Durafiber® wound dressings.
  • the present disclosure provides the ability to tailor mesh absorption, retention and shrinkage/expansion to the application at hand.
  • the small fiber sizes obtained also improves softness, conformability and leads to very high surface areas.
  • High surface area improves absorptive capabilities, hydration kinetics and drug release capabilities, among other properties.
  • the fibrous form factor allows for formation collection into novel form factors such as yarns, ropes, tubes, meshes, etc.
  • Example 5 Fibers with a poiyurethane core containing silver particles and a hydrophilic poiyurethane sheath.
  • thermoplastic poiyurethane material (HBPU) was used as the core material.
  • the electrospinning solutions were as follows: 4 wt% HLPU in TFE and 6 wt% HBPU in HFIP containing 30% silver nanoparticles with respect to the polymer.
  • the resulting fibers exhibited a core-sheath geometry in which silver was encapsulated and are shown in FIG. 21.
  • Silver is well-known for its antibacterial properties and such a mesh could be used for sustained release of silver for wound dressing applications.
  • other embodiments including incorporation of other particles and/or excipients into the core material to achieve different performance metrics.
  • cross-linked celluloses or other hydrophilic polymers can be incorporated into the core to further aid in the hydration properties of the resulting fiber.

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Abstract

Selon un aspect de l'invention, des fibres multicomposant sont fournies, qui comprennent (a) une âme polymère qui comprend un polymère de formation d'âme et (b) une gaine polymère qui comprend un polymère de formation de gaine qui est différent du polymère de formation d'âme. Des exemples de polymères de formation d'âme comprennent, par exemple, des polysiloxanes réticulés et des polymères thermoplastiques, parmi d'autres. Des exemples de polymères de formation de gaine comprennent, par exemple, des polymères solubles dans les solvants, des polymères dégradables et des polymères de formation d'hydrogel, parmi d'autres. D'autres aspects de la présente invention concernent des procédés de formation de telles fibres multicomposant. Par exemple, dans certains modes de réalisation préférés, des fibres multicomposant sont formées à l'aide de techniques d'électrofilage coaxial. Encore d'autres aspects de la présente invention concernent des mailles et d'autres articles qui sont formés à l'aide de fibres multicomposant.
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US10925998B2 (en) 2017-04-25 2021-02-23 Boston Scientific Scimed, Inc. Method of manufacturing a biocompatible composite material
EP3787554A4 (fr) * 2018-04-30 2021-06-16 Board of Regents of the University of Nebraska Endoprothèse-greffe
WO2019213161A1 (fr) * 2018-04-30 2019-11-07 Board Of Regents Of The University Of Nebraska Endoprothèse-greffe
CN114984295A (zh) * 2022-04-14 2022-09-02 广东云曌医疗科技有限公司 多孔纳米医用敷料及其制备方法
EP4269671A1 (fr) 2022-04-26 2023-11-01 EMPA Eidgenössische Materialprüfungs- und Forschungsanstalt Filage humide à base microfluidique de fibres polymères solides individuelles
WO2023209037A1 (fr) 2022-04-26 2023-11-02 Empa Eidgenössische Materialprüfungs- Und Forschungsanstalt Filage humide à base microfluidique de fibres polymères solides individuelles
CN115990293A (zh) * 2023-01-30 2023-04-21 博裕纤维科技（苏州）有限公司 医用多层自支撑载药壳-芯结构纳米纤维管的制备方法

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