JP2016519222A - Core-sheath fiber and method for making it and method for using it - Google Patents

Abstract

In accordance with one aspect of the present invention, multicomponent fibers are provided, the multicomponent fibers comprising: (a) a polymer core comprising a core-forming polymer; and (b) a polymer sheath comprising a sheath-forming polymer different from the core-forming polymer. Including. Examples of core-forming polymers include, for example, crosslinked polysiloxanes and thermoplastic polymers, among others. Examples of sheath-forming polymers include, for example, solvent-soluble polymers, degradable polymers, and hydrogel-forming polymers, among others. Another aspect of the invention relates to a method for forming such multicomponent fibers. For example, in certain preferred embodiments, the multicomponent fibers are formed using coaxial electrospinning techniques. Yet another aspect of the present invention relates to meshes and other articles formed using the multicomponent fibers.

Description

RELATED APPLICATION This application is related to US Provisional Application No. 61 / 852,224, filed March 15, 2013, which is the title of the invention “System and Method for Production of Silicone Fibers Using Coaxial Electrospinning”. And the benefit of US Provisional Application No. 61 / 861,629, filed Aug. 2, 2013, which is the title of the invention "Biocomponent Elastomer Hydrogel Fiber", each of which is hereby incorporated by reference in its entirety. Embedded in the book.

  This invention was made with government support under Grant No. 70NANB11H004, a technology innovation program grant awarded to the National Institute of Standards and Technology (NIST). The government has certain rights in the invention.

(Technical field)
The disclosure relates to, among other things, core-sheath fibers, methods for making core-sheath fibers, and devices and applications associated with core-sheath fibers.

  Fibers and collections of fibers have been used as materials in a variety of industrial applications, including medical and surgical applications ranging from sutures, especially wound dressings to skin grafts or artificial grafts. These applications are based on the unique properties of the fibers as material.

(Summary of the Invention)
According to one aspect of the invention, multicomponent fibers are provided, wherein the multicomponent fibers are (a) a polymer core comprising a core-forming polymer, and (b) at least partially surrounding the polymer core, the core-forming polymer. A polymer sheath comprising a different sheath-forming polymer. Examples of core-forming polymers include, for example, crosslinked polysiloxanes and thermoplastic polymers, among others. Examples of sheath-forming polymers include, for example, solvent-soluble polymers, degradable polymers, and hydrogel-forming polymers, among others.

  Another aspect of the invention relates to a method for forming such multicomponent fibers. For example, in various preferred embodiments, the multicomponent fibers are formed using coaxial electrospinning techniques.

  Yet another aspect of the present invention relates to meshes and other articles formed using the multicomponent fibers.

  These and many other aspects and embodiments of the present invention will be readily apparent to those of ordinary skill in the art after reviewing the following detailed description and claims.

FIG. 1 shows a photomicrograph of a cross section of a PLGA / PDMS sheath / core fiber formed in accordance with an embodiment of the present invention.

FIG. 2 shows the PDMS fiber of FIG. 1 after removal of the sheath layer.

3A-3B are top and cross-sectional photographs (before sheath removal (FIGS. 3A and 3C) and after sheath removal (FIGS. 3B and 3D) of a PLGA / PDMS sheath / core fiber formed in accordance with one embodiment of the present invention. Both).

FIG. 4 shows images of water drops (left) and oil drops (right) placed on a PDMS mesh according to the present invention.

FIG. 5 is a stress-strain diagram illustrating the mechanical properties of a PDMS mesh according to the present invention compared to a cast PDMS film.

FIGS. 6A-6B show photomicrographs of cross sections of three different sheaths: PLGA / PDMS sheath / core fibers electrospun at different core flow rates in accordance with one embodiment of the present invention.

FIG. 7 shows a photomicrograph of a PVP / PDMS sheath / core fiber formed according to one embodiment of the present invention. These are: (A) Cross-section of core-sheath fibers with PVP cured at 100 ° C; (B) Same fibers as in (A) after undergoing water extraction; (C) PVP cured at 150 ° C Cross section of core-sheath fiber; (D) shows the same fiber as in (C) after undergoing water extraction.

FIG. 8 shows the FTIR (Fourier) of a pure PDMS film, a pure PVP film, and a PVP / PDMS sheath / core fiber formed according to one embodiment of the invention (cured at 100 ° C.) when dried and wet. A conversion infrared spectroscopy) scan is shown.

FIG. 9 shows an FTIR scan of pure PDMS film, pure PVP film and PVP / PDMS sheath / core fibers formed according to one embodiment of the invention (cured at 150 ° C.) when dried and wet. Show.

FIG. 10 is a stress diagram illustrating the mechanical properties of a PVP / PDMS sheath / core fiber formed according to one embodiment of the invention (cured at 100 ° C. and 150 ° C.) when dried and wet. FIG.

FIGS. 11A and 11B illustrate a balloon formed from a hydrated PVP-PDMS fiber mesh cured at 100 ° C. at two expansion levels according to one embodiment of the present invention.

FIG. 12 illustrates a fiber having an HLPU sheath and an HBPU core formed herein with four hydrophilic polyurethane (HLPU): more hydrophobic polyurethane (HBPU) ratios according to various embodiments of the present invention (herein A photomicrograph of HLPU / HBPU sheath / core fiber).

FIG. 13 shows the swelling and tensile strength as a function of HLPU content of a mesh formed from HLPU / HBPU sheath / core fibers formed according to various embodiments of the present invention.

FIG. 14 shows swelling and shrinkage as a function of HLPU content of a mesh formed from HLPU / HBPU sheath / core fibers formed in accordance with various embodiments of the invention.

FIG. 15 shows the swelling of a mesh formed from four different HLPU / HBPU sheath / core fibers (formulations AD) formed in accordance with the present invention, as well as two commercial wound dressings.

FIG. 16 shows the wet tensile strength of a mesh formed from four different HLPU / HBPU sheath / core fibers (formulations AD) formed in accordance with the present invention, as well as two commercial wound dressings.

FIG. 17 shows the shrinkage of a mesh formed from four different HLPU / HBPU sheath / core fibers (formulations AD) formed in accordance with the present invention, as well as two commercial wound dressings.

18A and 18B show photomicrographs of meshes formed from HLPU / HBPU sheath / core fibers before and after annealing, respectively, according to one embodiment of the present invention.

FIG. 19 shows a mesh formed from annealed (B annealed) and non-annealed (B standard) HLPU / HBPU sheath / core fibers formed in accordance with the present invention, and phosphorus from two commercial wound dressings. Acid buffered saline (PBS) retention is shown.

FIG. 20 shows a mesh formed from annealed (B annealed) and unannealed (B standard) HLPU / HBPU sheath / core fibers formed in accordance with the present invention, and shrinkage of two commercial wound dressings. / Indicates expansion.

FIG. 21 is a photomicrograph of HLPU / HBPU sheath / core fibers with encapsulated silver nanoparticles.

Detailed Description of Preferred Embodiments)
According to one aspect of the disclosure, a multicomponent fiber is provided that includes a polymer core and a polymer sheath that at least partially surrounds (ie, encloses) the core.

  As used herein, “fibers”, “microfibers” and “nanofibers” are used synonymously to refer to elongated structures that differ only in size (“microfiber” is a number of fibers. “Nanofiber” indicates that the fiber has a cross-sectional diameter on the order of several nanometers to several hundred nanometers, and “fiber” indicates any size Also shows fibers).

  The fibers according to the present disclosure can thus be formed in a wide variety of sizes. Preferred overall fiber diameters are 0.05 to 50 microns (μm) (eg 0.05 to 0.1 microns to 0.25 to 0.5 microns to 1 to 2.5 microns to 5 to 10 microns to More preferably in the range of 0.1 to 20 microns, among the possible dimensions. Preferred core diameters are 0.01-10 microns (e.g., 0.01-0.025 microns to 0.05-0.1 microns to 0.25-0.5 microns to 1), among other possible dimensions. Range from ˜2.5 microns to 5-10 microns). A preferred sheath thickness is 0.02 to 25 microns (eg, 0.02 to 0.05 microns to 0.1 to 0.25 microns to 0.5 to 1 microns to 2.5 to 5 microns to 10 to 25 microns). And more preferably, among the possible dimensions, it ranges from 0.2 to 18 microns.

  The ratio of sheath volume to core volume can vary widely. Preferred sheath volume: core volume ratios range, for example, from 100: 1 to 1: 100, and, among other values, for example, 100: 1 to 50: 1 to 25: 1 to 10: 1 to 5: 1. It ranges from 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 according to the present disclosure can be formed using a variety of fiber spinning techniques, including various melt spinning methods and solvent spinning methods. Accordingly, although solvent spinning techniques, and more specifically, electrostatic solvent spinning techniques are described in detail herein, the present invention is not limited to such techniques. Further exemplary techniques for forming multicomponent fibers include, among others, hot melt spinning, melt electrospinning, centrifugal fiber spinning, wet spinning, dry spinning, gel spinning, gravity spinning, extrusion, Examples include extrusion spinning, and rapid prototyping. Using these and other techniques, (a) a polymer core that includes a core-forming polymer, and (b) a polymer sheath that at least partially surrounds the polymer core and includes a sheath-forming polymer different from the core-forming polymer. Including multicomponent fibers can be formed.

  Electrospinning is a process that uses charge to draw very fine, typically microscale or nanoscale, fibers from a liquid. Solvent electrospinning utilizes the force of electricity applied to the polymer solution to induce an electrospinning jet. As the stream associated with the jet moves into the air (or other atmosphere), the evaporation of the solvent results in a single long polymer fiber deposited on the grounded collector. The collected fibers can result in the formation of meshes that can be used in various technologies in the medical and non-medical industries (eg, drug delivery devices, tissue engineering, nanoscale sensors, wound dressings, self-healing coatings, among others) self-healing coating) and filters.

  As used herein, a “mesh” is a structure formed by a collection of one or more fibers that are intertwined to form a three-dimensional network. The mesh includes a woven mesh or a non-woven mesh.

  Meshes according to the present disclosure can vary widely in thickness, with preferred thicknesses ranging from 10 to 5000 microns (e.g., 10 to 25 microns to 50 to 100 microns to 250 to 500 microns to 1000 to 2500 microns to 5000, among other values). Can range in the micron range).

  Meshes according to the present disclosure can vary widely in porosity. In certain embodiments, meshes of the present disclosure have a porosity of 99% or less (eg, 99% -90% -80% -70% -60% -50% -40% -30% -20% -10% % Range). The porosity can be measured by determining the volume of the polymer and dividing that amount by the volume of the mesh. In this regard, polymer volume = mesh mass ÷ polymer density; mesh volume = mesh length × mesh width × mesh thickness = mesh area × mesh thickness; and mesh porosity = (mesh volume−polymer volume) ÷ mesh volume. In various embodiments, the porosity of a given mesh can be reduced by annealing the mesh at a temperature and for a time over which a decrease in mesh porosity is observed.

(Electrospinning)
Conventionally, core-sheath electrospinning (also referred to herein as coaxial electrospinning) utilizes two concentric needles to provide two types of solutions (specifically, an inner core polymer solution and an outer sheath polymer). Solution) is delivered separately. The core solution is delivered through the inner needle, while the sheath solution is delivered through the outer needle. Upon actuation of the electric field, the two different polymer solutions are injected in a continuous stream toward the grounded collector; this configuration forms a single core-sheath Taylor cone at the needle tip, and the core-sheath Leading to the formation of fibers. However, the generation of core-sheath fibers using needles has limited throughput.

In certain embodiments, the core-sheath fiber is a high-throughput core-sheath needle-less electrospinning fixture (which includes one or more slits on the surface of a hollow container). Utilized to coexist many materials at multiple sites forming a Taylor cone, thereby promoting the formation of multiple electrospun jets and thus multiple electrospun fibers) . Thus, slits in the surface of the hollow container can result in high-throughput generation of core-sheath fibers. For further information, see, for example, US Patent Publication No. 2012/0193936 (Sharma et al.) And US Patent Publication No. 2013/0241115 (Sharma et al.), The disclosures of which are incorporated by reference. .

  In electrospinning, each jet so formed results in one long continuous fiber being collected. In a typical operation of the needle-free installation, about 10 jets are formed along the length of the slit; this collected mesh is therefore from about 10 very long, intertwined fibers. Composed. In contrast, during the operation of the open bath free surface electrospinning used in the high-throughput core-sheath and needleless electrospinning equipment, hundreds of jets are formed and the drum Disappears with each rotation. Thus, the resulting mesh consists of thousands of relatively short fibers.

  The design of the needle-free electrospinning facility takes into account processing parameters that may allow greater control over fiber diameter. For example, in addition to solution properties, solution flow rate can be manipulated to control fiber diameter. In addition, the number of jets generated can be controlled, which can lead to fiber diameter differences.

  The fibers of any embodiment of the present disclosure can therefore be collected in the form of a nonwoven mesh. However, alternative embodiments may be used as a conditioned fiber (when passing through a gap alignment or rotating drum), twisted yarn, rope, some pattern, or any other method of fiber collection known in the field of electrospinning. , Including collected fibers.

(Fiber with silicone component)
Various aspects of the invention relate to multicomponent fibers formed using silicone polymers (also referred to herein as “silicones”, “siloxane polymers” or “polysiloxanes”). For example, in certain embodiments, (a) a polymer core that includes one or more silicone polymers, and (b) a polymer sheath that at least partially encloses the core and includes one or more additional polymers other than silicone. Containing multicomponent fibers (and vice versa) are formed.

The present disclosure provides a plurality of siloxane units:
Applicable to all siloxanes, including polysiloxanes (ie, compounds having —Si—O—Si bonds),
Where R ₁ and R ₂ are organic radicals, for example, linear, branched or cyclic alkyl groups which may be substituted or unsubstituted (eg methyl, ethyl, propyl Groups, isopropyl groups, butyl groups, isobutyl groups, sec-butyl groups, tert-butyl groups, cyclohexyl groups, etc.) and substituted or unsubstituted aryl groups (eg phenyl groups, p-, m- or o-alkyl substituted phenyl groups, etc.). R ₁ and R ₂ may be the same or different.

  In various embodiments, polysiloxanes included as PDMS can be functionalized by various mechanisms (eg, plasma, UV, CVD, etc.) to modify surface properties (eg, hydrophobicity, etc.) or specific chemistry. Interaction (eg, antibody binding) may be provided. The fibers can be functionalized to produce immobilized biomolecules on the surface and / or in the bulk. Functionalization can provide materials with many new properties, including biological effects, sensor applications. Microfibers and nanofibers further enhance these benefits, for example, by providing high surface areas and small pores.

  In this regard, functional groups that are polymerized as pendant groups attached to the siloxane (eg, hydride, hydroxyl, amine, isocyanate, epoxy, etc.) add chemical activity and versatility, and, among other properties, mechanical Can be used to modify mechanical properties, swelling and solvent resistance, and refractive index. Coaxial electrospinning of polysiloxanes as described herein can be combined with functionalization to obtain silicone microfibers and nanofibers with various properties that make them useful in further applications. For example, the process of making the fibers more hydrophilic provides elastic, durable fibers that wet more easily. In some embodiments, the functionalizing portion of the PDMS is incorporated into the fiber. Upon curing, functional moieties in the fibers are incorporated into the PDMS through siloxane chemistry. This allows for one-step functionalization of the PDMS. In one specific embodiment, the PDMS surface is prepared according to Bo Huang et al. , “Phospholipid biotinylation of polydimethylsiloxane (PDMS) for protein immobilization” as described in Lab Chip, 2006, 6, 369-373. Can be functionalized with These biotin groups can then be further modified with avidin conjugated to the species of interest (eg, protein, antibody or fragment thereof) to functionalize the silicone surface. This is useful, for example, in removing proteins from lipids (eg, protein separation) or in medical implants where preferential binding of certain proteins is advantageous (eg, improved endothelial cell interaction). It can be.

  Several classes of polymers, including various siloxane polymers, have been shown to be difficult to electrospin due to their low molecular weight and fluidity. In this regard, various polysiloxanes remain fluid until they are cross-linked, which does not allow for sufficient polymer chain entanglement to form fibers.

  For example, polydimethylsiloxane (PDMS) is a silicon-based organic polymer that belongs to a larger group of siloxane polymers as described above that generally exhibits properties of elasticity and durability. The ability to control fiber diameter, as well as the ability to form fibers and constructs made from PDMS and other siloxane polymers exhibiting such properties, as well as a variety of other applications, is very important in medical technology. Is advantageous. Attempts have been made to electrospin PDMS fibers, but the techniques that have been developed so far use blended polymer systems (ie, not pure PDMS), so that pure PDMS fiber constructs (eg, There is currently no electrospinning method known to the inventors for producing a mesh).

  Accordingly, in some aspects of the present disclosure, core-sheath electrospinning techniques are provided. This can be used to form fibers containing silicone material that was not previously electrospun using known techniques. The fibers formed by the techniques described herein include a silicone material as the core material and a different polymer material as the sheath material. After fiber formation and / or collection, the core-sheath fibers are typically crosslinked by a suitable mechanism. For example, the fibers can be cured at room temperature overnight, or at temperatures up to 100 ° C. for several hours, among other crosslinking techniques.

  In certain embodiments, the polymer sheath can be formed from a hydrophilic or hydrogel material, which will be discussed in more detail below.

  In certain embodiments, the polymer sheath can be formed from a material 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).

  Examples of degradable polymers include, among others, one or more of the following: (a) polyglycolide (PGA) (also referred to as polyglycolic acid), polylactide (PLA) (also referred to as polylactic acid) Polyester homopolymers and copolymers (especially poly-L-lactide, poly-D-lactide and poly-D, L-lactide, poly (lactide-co-glycolide) (PLGA), polycaprolactone, polyvalerolactone, poly (β -Hydroxybutyrate), polygluconate (including poly-D-gluconate, poly-L-gluconate, poly-D, L-gluconate), poly (p-dioxanone), poly (lactide-co-δ-valero) Lactone), poly (lactide-co-ε-caprolactone), poly (lactide) co-β-malic acid), poly (β-hydroxybutyrate-co-β-hydroxyvalerate)), (b) polycarbonate homopolymers and copolymers (eg, poly (tymethylene carbonate), poly, among others) (Lactide-co-trimethyl carbonate) and poly (glycolide-co-trimethylene carbonate)), (c) poly (orthoester) homopolymers and copolymers (for example, synthesized by copolymerization of various diketene acetals and diols, among others). ), (D) polyanhydride homopolymers and copolymers (eg, inter alia, poly (adipic anhydride), poly (suberic anhydride), poly (sebacic anhydride), poly (dodecanedioic anhydride) ), Poly (maleic anhydride) and poly [1,3-bis (p- carboxyphenoxy) methane anhydride]), (e) polyphosphazene (for example, among others, polyphosphazenes were aminated and alkoxy substituted), and (f) amino acid based polymers.

  Examples of water-soluble polymers include non-crosslinked hydrophilic polymers, which can 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 its salt, methacrylic acid and its salt, hydroxyethyl methacrylate, acrylamide, N, N-dimethylacrylamide, N-hydroxymethylacrylamide, alkyloxazoline, saccharide monomer ( For example, polysaccharides such as dextran, alginate, etc.), and amino acids (eg, hydrophilic polypeptides and proteins (eg, gelatin, etc.)). When crosslinked, the hydrophilic polymers described above are useful as hydrogels.

  For standard nonwoven materials, the microstructure is highly dependent on the fiber diameter. Thus, the advantage of this core-sheath manufacturing process where the sheath is subsequently removed is the ability to obtain pore size, porosity and other microstructural features. Using the above high-throughput core-sheath and needle-free electrospinning equipment (see, for example, US Patent Publication Nos. 2012/0193936 and 2013/0241115 (Sharma et al.)), The sheath to core thickness ratio is It can vary to provide larger pore sizes and thinner fibers, or higher porosity and thinner fibers than can be obtained with other fabrication techniques.

(Fibers with hydrogel components and variable hydrophilic / hydrophobic components)
Various aspects of the invention relate to multicomponent fibers formed using hydrogels. For example, in certain embodiments, a multi-component fiber is formed that includes (a) a polymer core that includes one or more core-forming polymers, and (b) a polymer sheath that includes one or more hydrophilic or hydrogel-forming polymers. Is done.

  Various aspects of the present invention include: (a) a polymer sheath comprising one or more hydrophilic polymers; and (b) a polymer core comprising one or more polymers that are more hydrophobic than the one or more hydrophilic polymers. Containing multicomponent fibers. Conversely, another aspect of the present invention includes (a) a polymer core comprising one or more hydrophilic polymers, and (b) one or more polymers that are more hydrophobic than the one or more hydrophilic polymers. It relates to multicomponent fibers comprising a polymer sheath.

  Polymers for use as core and / or sheath polymers include (wet weight-dry weight) / dry weight when immersed in an aqueous medium (eg, water, PBS, etc.) at 25 ° C. for 1 hour. Water absorption values, calculated as (× 100), generally ranging from 0 to 1000% or more of water (eg 0% to 1% to 2.5% to 5% to 10% to 25% to 50% to In the range of 100% to 250% to 500% to 1000% or more). As used herein, a “hydrophilic polymer” is one having a water absorption value ranging from 5 to 1000% water. A “more hydrophobic polymer” (also referred to herein as a “less hydrophilic polymer”) is defined as a polymer that absorbs less water than the given polymer being compared.

  In some embodiments, the core and sheath polymer have a ratio of sheath polymer water absorption value: core polymer water absorption value ranging from 2: 1 to 100: 1 (eg, 2: 1 to 5: 1 to 10: 1). : Ranging from 1 to 20: 1 to 50: 1 to 100: 1), among other possible values, preferably in a range of 5: 1 to 20: 1 in certain embodiments, Selected. By way of illustration, 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%, resulting in a sheath: core water absorption ratio of 10: 1. .

  Hydrogels contain a three-dimensional crosslinked network of hydrophilic polymers that have the ability to absorb substantial amounts of water. Hydrogels have long been used in many applications in the medical field, ranging from drug delivery to tissue engineering scaffolds. Despite many potential applications, hydrogels have been limited in healthcare and other fields due to lack of structural control and poor understanding of the mechanical properties of hydrogels. . Others in this field have studied to strengthen hydrogels with various additives. Still others reinforce the hydrogel by making polymer fibers or polymer fiber constructs (eg, meshes) and then immersing them in a hydrogel or hydrogel-forming polymer prior to crosslinking the polymer. Aimed at. Since such methods and structures are generally not effective, there is still a need for a hydrogel structure with the desired properties.

  In certain aspects of the present disclosure, electrospinning is used to form a fiber core that includes one or more fibrogenic polymers that are at least partially surrounded by a sheath that includes one or more hydrogel-forming polymers. . The resulting composite fiber is subjected to a crosslinking step (eg, by application of energy such as heat, visible light or ultraviolet light, by application of a crosslinking agent, etc.) as necessary to form a hydrogel-forming polymer, a core The polymer or both can be cross-linked. The result is a composite fiber with mechanical and hydration properties that differ from either material alone. These composite fibers can be collected, formed, or processed into various shapes (eg, tubes, meshes, threads, etc.) for use as medical devices or other products.

Polyurethane can be used as the core and / or sheath polymer in various embodiments. Polyurethanes are generally formed from diisocyanates and long chain diols, and typically chain extenders. Aromatic diisocyanates may be selected from among the following suitable members, among others: methylene diphenyl diisocyanate (MDI), toluene diisocyanate (TDI), naphthalene diisocyanate (NDI), para-phenylene diisocyanate (PPDI), 3,3′-tolidine. (Tolidene) -4,4′-diisocyanate and 3,3′-dimethyl-diphenylmethane-4,4′-diisocyanate. Non-aromatic (aliphatic) diisocyanates may be selected from among the appropriate members, among others: hexamethylene diisocyanate (HDI), dicyclohexylmethane diisocyanate (H ₁₂ MDI), isophorone diisocyanate (IPDI), cyclohexane diisocyanate (CHDI), among others. ), 2,2,4-trimethyl-1,6-hexamethylene diisocyanate (TMDI), and meta-tetramethylxylene diisocyanate (TMXDI). Long chain diols include polyether diols (eg, polyethylene glycol, polyoxypropylene glycol, polytetramethylene ether glycol, etc.), polyester diols (eg, polybutanediol adipate, polyethylene adipate, polycaprolactone diol, etc.), and polycarbonate diols Is mentioned. Other long chain diols include diols such as 1,4-butanediol, among others.

  Polyurethanes other than those described in the previous paragraph may also be used as the core and / or sheath polymer in various embodiments.

  Hydrogels for use in the present disclosure are hydrophilic that are cross-linked through appropriate mechanisms (eg, covalent and / or non-covalent cross-linking (eg, by ionic cross-linking, physical cross-linking, etc.)). Includes those formed from polymers.

  Examples of hydrophilic polymers that can be crosslinked include various hydrophilic polymers such as those shown above. Further examples of hydrophilic polymers include hydrophilic polyurethanes (eg, polyurethanes having hydrophilic segments), which can be physically cross-linked (eg, via hard segments present in the polyurethane). Specific hydrophilic polyurethanes include, among others, aliphatic polyether-based polyurethanes and aromatic polyether-based polyurethanes. It is further noted that the hydrophilic polymers shown above can be used as hydrophilic segments in polyurethanes in certain embodiments.

  Examples of core-forming polymers that in many embodiments include thermoplastic polymers and various hydrophilic / hydrophobic polymers include, among others, silicones (polysiloxanes) such as those described above, thermoplastic polyurethanes (eg, In particular, aliphatic polyether-based polyurethanes and aromatic polyether-based polyurethanes) and polyamides (eg, nylon-6,6, nylon-6, nylon-6,9, nylon-6,10, nylon-6, 12, nylon-11, nylon-12, nylon-4, 6, etc.). Examples of core-forming polymers further include, among other things, homopolymers and copolymers (including block copolymers) comprising one or more of the following monomers: (a) unsaturated hydrocarbon monomers (eg, ethylene, propylene) , Isobutylene, 1-butene, 4-methylpentene, 1-octene and other α-olefins, isoprene, butadiene, etc.); (b) halogenated unsaturated hydrocarbon monomers (eg, tetrafluoroethylene, vinylidene chloride, vinylidene fluoride) (C) unsubstituted vinyl aromatic monomers (eg, styrene, 2-vinylnaphthalene, etc.) and vinyl substituted aromatic monomers (eg, α-methyl), chlorobutadiene, vinyl chloride, vinyl fluoride, etc.); Styrene), ring-substituted bi Vinyl aromatic monomers, including nyl aromatic monomers; and (d) relatively hydrophobic (meth) acrylic monomers (alkyl (meth) acrylates (eg, isopropyl acrylate, butyl acrylate, sec-butyl acrylate, isobutyl acrylate, cyclohexyl) Acrylate, tert-butyl acrylate, hexyl acrylate, 2-ethylhexyl acrylate, dodecyl acrylate, hexadecyl acrylate, and isobornyl acrylate, isopropyl methacrylate, isobutyl methacrylate, t-butyl methacrylate, cyclohexyl methacrylate, 2-ethylhexyl methacrylate, octyl methacrylate, Dodecyl methacrylate, hexadecyl methacrylate, octadecyl methacrylate Relates, isobornyl methacrylate, etc.), arylalkyl (meth) acrylates (eg, benzyl acrylate, benzyl methacrylate, etc.), and halo-alkyl (meth) acrylates (eg, 2,2,2-trifluoroethyl acrylate). ).

  The advantages associated with providing a multi-component fiber with a hydrogel sheath and a different core material from the sheath material are combined with good water absorption and retention properties (as a result of the hydrogel material), together with the desired mechanical properties (e.g. Fibers, meshes and other constructs having strength, elasticity, durability and shrinkage (as a result of the core material) can be formed.

(Fiber with silicone core and removable sheath)
As previously indicated, certain aspects of the present disclosure include: (a) a polymer core that includes one or more silicone polymers, and (b) a polymer sheath that includes one or more additional polymers other than silicone. Concerning component fibers. In certain embodiments, the polymer sheath can be dissolved from the silicone core, can be degraded or otherwise removed, and can be formed from a material that later leaves pure silicone fibers. Examples of such materials include, inter alia, degradable polymers and solvent-soluble polymers such as those shown above, including water-soluble polymers. As elsewhere herein, the fibers can be formed or processed into various shapes (eg, tubes, meshes, threads) for use as medical devices or other products.

  In some embodiments, the silicon core-forming polymer is co-electrospun with a removable (eg, dissolvable or degradable) sheath-forming polymer to create a new composite fiber. The electrospinning can be accomplished by needle-free electrospinning, coaxial electrospinning, slit surface electrospinning, or any other suitable technique known in the field of fiber spinning.

  In one preferred embodiment, detailed in Example 1 and Example 2 below, the 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 a prepolymer and a cross-linking agent that initiates the cross-linking reaction (exposure to heat promotes this reaction). As used herein, a “prepolymer” is a polymeric material that is subjected to a crosslinking process or other curing process to create a crosslinked polymer. In other embodiments, the two-component PDMS system can be cured by exposure to UV radiation. In still other embodiments, the two-component PDMS system can be cross-linked to an elastic body through free radical, condensation or addition reactions. Alternatively, a one-component PDMS system that cures upon exposure to atmospheric moisture or upon photocuring can be used, among other techniques. Any of these variations in PDMS chemistry, or other polymers that require physical or chemical cross-linking to become a solid, can be used in the fibers and methods described herein.

  Thus, although polysiloxane (ie, PDMS) is illustrated as a preferred embodiment, other embodiments may use polymers that require cross-linking to become solids (eg, thermosetting polymers). Examples include other polysiloxanes and certain types of polyesters, polyurethanes, polyimides, epoxides, and the like.

  While degradable polymers (ie, poly (lactide-co-glycolide)) are illustrated as preferred embodiments, other embodiments may include other degradable polymers or solvent-soluble polymer sheaths (eg, non-crosslinked PEO). , Formed from a water-soluble sheath material such as PVA, gelatin, dextran, carbohydrates, etc., which can then be removed by dissolution. Embodiments using an aqueous solvent as a solubilizer generally do not cause PDMS fiber swelling.

  In some embodiments, the sheath is etched away using acid.

  Depending on the mechanical properties of the sheath polymer, mechanical failure can be used to remove the sheath. Any combination of the above methods, or other suitable means, can be used to remove the sheath from the underlying core.

  In some embodiments, such as small molecule drugs, anesthetics, procoagulants, anticoagulants, antimicrobials, biologics, RNAi, genetic material, gene vectors, vaccines, or particles (eg, silver nanoparticles) The therapeutic agent is in the polysiloxane core.

  In some embodiments, a porogen (eg, selected from salts, sugars, etc.) is incorporated into the polysiloxane core. Upon subsequent sheath removal, the porogen is also removed. This leaves behind fibers with porous or rough surface features that can improve hydrophobicity, among other properties. Alternatively, the porogen can be incorporated into the sheath such that a certain percentage of the porogen is located at the core and sheath interface after fibril formation. When the sheath is removed, a negative imprint of porogen is present on the surface of the polysiloxane fiber. The surface of the polysiloxane fibers can also be roughened by a suitable etching process (eg laser etching) or mechanical means.

  Further, porosity can be introduced into the fibers of the present disclosure as the product of a cross-linking reaction that forms the fibers. For example, isocyanate functionalized PDMS can react with water to form a porous foam fiber. Another example is acetoxy-functionalized PDMS blended with sodium bicarbonate. The acetic acid byproduct of the cross-linking reaction can react with sodium bicarbonate to produce gas and porosity, thus allowing the formation of porous foam fibers.

  Manipulation of fiber size can produce a variety of fiber properties. For example, in filtration applications, finer fibers with large pores or higher porosity can increase permeability and surface area. Polysiloxane materials (eg, PDMS) as described herein provide small pore size and high permeability combined with high durability, thermal and oxidative stability, and flexibility. Furthermore, the polysiloxane fiber mesh formed in accordance with the present disclosure has a high surface area due to small size fibers. This can promote adhesion and wetting if desired. In some embodiments, these same properties may be useful in medical applications where cell infiltration into the fiber is desired. In particular, thinner fiber diameters generally promote cell interactions, inward growth and proliferation, while larger pores and higher permeability generally promote nutrient, cytokine and gas exchange while simultaneously Improves movement as well.

  In a further embodiment, the sheath is left on the core to form a composite fiber that includes a PDMS core with unique properties and a polymer sheath (eg, nylon, polyethylene, polystyrene, polycarbonate, etc.). In some embodiments, when immersed in water, the outer sheath may form a hydrogel to fill the porosity of the PDMS fiber mesh.

  In a further embodiment, the core and sheath polymer are reversed and polysiloxane is used as the sheath polymer that coats the core polymer. This allows the formation of bicomponent fibers with polysiloxane on the outside. Furthermore, removal of the core polymer results in polysiloxane hollow fibers.

  Small diameter fiber meshes can provide higher surface area, higher permeability and lower pore size than meshes made from larger diameter fibers. The present disclosure thus provides a material that combines the benefits of a polysiloxane such as PDMS and the fine diameter fiber mesh. For example, solvent resistant filters or elastic biocompatible microfiber or nanofiber medical device components (eg, heart valve leaflets, vascular grafts, stent graft coatings) may be formed.

  In this regard, in some embodiments, a silicone mesh may be used in the heart valve leaflets. Replacement heart valves in some cases use synthetic materials to recreate the native leaflet. Natural valve membranes are thin, very flexible and durable. In addition to these properties, the valve membrane must be non-thrombogenic. Promoting endothelialization is one of the best ways to provide non-thrombogenic implants. The microfiber structure provided by electrospun silicone is believed to promote endothelial cell proliferation. However, this same porosity can lead to blood passing through the pores of the mesh and reduced blood flow control by the valve. This phenomenon, however, is expected to be transient as proteins and cells are trapped in the pores. In a preferred embodiment, the silicone microfiber mesh is electrospun to a thickness of between 100 and 1000 microns (μm). The target fiber diameter is between 500 nm and 10 μm. These meshes are then cut into the appropriate shape and attached to the body, which is implanted via laparotomy or minimally invasive surgery. Alternative embodiments include: a membrane (either sandwiched on one side of the mesh or between two meshes to prevent blood flow through the mesh) For example, providing silicone, PLGA); functionalizing silicone with proteins or antibodies (eg, CD34, VEGF) to promote tissue ingrowth and re-endothelialization; frames (eg, polymer fibers To help form the valve membrane and / or provide attachment to the body; to a biocompatible fibrous structure that creates a composite implant Electrospinning (eg, the fibers provide additional mechanical strength or varying stiffness across the valve membrane); and the fibers Coating or as a functionalized, reducing the thrombogenicity (e.g., heparin).

  In some embodiments, the silicone mesh is stent graft coated in a manner similar to a heart valve membrane, except that the silicone fibers are electrospun onto a tubular collector to form a silicone microfiber or nanofiber tube. Can be used. A preferred mesh thickness is between 100 and 1000 microns. The target fiber diameter is between 500 nm and 10 μm. This tube can then be attached to a stent to form a stent graft. Alternatively, the fibers can be electrospun directly onto the stent. The alternative embodiment described with respect to the heart valve is also applicable here. Advantages include the fact that silicone microfibers and / or nanofibers promote cell ingrowth while providing an elastic, biocompatible, durable implant. In some embodiments, silicone mesh can be used in vascular grafts similar to stent graft designs, except that the tube is not attached to the stent and the preferred mesh thickness range is larger (100-5000 microns).

  In some embodiments, silicone mesh can be used in bioengineered blood vessels. Very similar to the vascular graft described above, the silicone mesh can be shaped into a tube and cells can be seeded ex vivo. These cells (typically fibroblasts, smooth muscle cells, and endothelial cells) may have various conditions (eg, pulsatile flow, constant flow, or Incubate in a nutrient rich environment under no flow). The silicone microfibers and nanofibers provide advantages in promoting cell infiltration and proliferation, as well as providing typical elastic properties to blood vessels. The silicone tube can be used alone or in combination with other natural materials (eg, collagen) or synthetic materials (eg, PTFE, ePTFE, polyurethane). In other embodiments, the graft is seeded with cells and transplanted without meaningful incubation or transplanted without cell seeding. In the latter case, host-derived cells infiltrate and settle in the graft.

  In some embodiments, silicone mesh can 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 mesh provides a strong set of mechanical properties and also promotes cell ingrowth. The elasticity, durability, biocompatibility and low thrombogenicity of silicones improve the performance of these implants. In one embodiment, the silicone microfiber or nanofiber mesh is shaped and implanted into a tube. The tube can be pretreated by functionalization or coating with other materials (eg, heparin, collagen, gelatin, growth factors) to improve integrity or cell ingrowth. In other embodiments, the silicone mesh may be combined with other natural or synthetic materials as a sheet or mesh to form a composite layered structure. This layered structure may improve the mechanical properties, inclusion of blood immediately after implantation, or long-term durability or performance.

  In some embodiments, silicone fibers electrospun into a flat mesh configuration having a thickness of 500-5000 microns can be used in hernia mesh. In order to improve the mechanical properties, composites can be formed with these fibers by electrospinning directly onto biocompatible polymer fibers in the desired configuration. These fibers can also be provided in a configuration that improves the suture performance of the mesh. In alternative embodiments, the mesh may be functionalized to improve tissue ingrowth or integration using the various methods described above.

  In some embodiments, the silicone mesh can be used in a double coating. In neurosurgical procedures where the dura is injured, it is desirable to provide a coating to reclose the membrane. Silicone microfiber or nanofiber mesh (optionally combined with a polymer membrane (eg, silicone, PLGA, collagen) may be used for this purpose).

  In other embodiments, the silicone mesh can be used in wound dressings. The disadvantages of wound dressings include adhesion to the wound and permeability to air and moisture (wound exudate). In one embodiment, the silicone is electrospun into a mesh between 100 and 5000 microns thick. A preferred fiber diameter is between 500 nm and 10 microns. The electrospun silicone is non-adhesive to the wound, provides high permeability and is used as a wound contact layer in a bandage. In another embodiment, the silicone bandage may be supplied separately and medical staff may place additional gauze or other bandages in layers over the silicone bandage. In another embodiment, the silicone mesh is combined with gauze or other backing material as part of the final product to absorb fluid and protect the wound. In still other embodiments, the silicone can be made with a therapeutic agent (eg, antibiotics, antifungal agents, topical analgesics, antiseptics (eg, iodine), etc.). Another embodiment provides a silicone mesh that is treated with or manufactured with a hydrogel sheath (eg, PEG) to provide moisture to the wound bed. The advantage of the silicone mesh in this case is that the high porosity can include the hydrogel material while at the same time assisting in the removal when the bandage needs to be removed. In yet other embodiments, the silicone mesh is fabricated for use in negative pressure closure therapy. In this case, the mesh is sized to be compatible with these devices and placed in the wound environment as negative pressure is applied. This high permeability and porosity allows exudate removal and non-adhesive bandages when the bandage must be removed. For application in negative pressure closure therapy, the silicone fibers can be electrospun onto the collector with a shape and topography similar to the intended treatment site (eg, face, head, etc.). In this way, the dressing may improve the therapy by an improved fit to the wounded tissue.

  In some embodiments, silicone mesh can be used in hemostatic applications. For hemostatic applications, the device is structured much like a wound dressing, but the silicone microfiber or nanofiber is a thrombus or coagulant (eg thrombin, kaolin, chitosan, fibrin, etc.) ) Or surface modification. The silicone provides a non-adhesive bandage that can be easily removed. In addition, its high permeability and porosity allows blood to penetrate and a significant amount of surface area to come into contact with the thrombus formation promoter. This open structure also allows clotting factors to diffuse back into the wound and promote clot formation. This material can also be integrated as a non-adhesive layer on other bandages (eg, combat gauze); for this application, the fibers may be made with a prothrombotic or procoagulant. And it doesn't have to be.

  In other embodiments, the silicone mesh can be used in filtration applications. The silicone mesh can be used as a filter for air, other gases, liquids, slurries or particles or as part of a filter. Its high solvent resistance and durability offers advantages over other microfiber and nanofiber filters. In particular, its low pore size and high permeability of electrospun microfiber or nanofiber mesh is desirable for filters. In addition, its elastic nature provides a way to clean the filter. Simply extending the material biaxially increases its pore size from the surroundings or otherwise. A gas or liquid back flow then provides a way to clean the pores from debris or other materials. In a similar manner, the cake formed on the intake side of the filter can be easily removed by extending the silicone mesh and dropping the cake. The silicone microfiber or nanofiber mesh can be used alone (preferred thickness is 100 microns to 1 cm). Alternatively, the silicone mesh can be constructed as part of a layered filter using other commercially available filter materials. In this case, the silicone can be electrospun directly onto another material, placed on the other material during assembly, or a large opening to provide wire or mechanical support. It can be electrospun onto other fiber mesh that it has.

  In some embodiments, silicone mesh can be used in drug delivery. Drugs can be incorporated into the silicone microfibers or nanofibers for delivery to a patient. In one embodiment, the silicone mesh is formed of a drug in the silicone solution and placed on the skin for dermal or transdermal delivery. In another embodiment, the silicone microfiber or nanofiber is formed into a mesh, tube or other structure and implanted to deliver the drug therein. This includes the mouth or other body openings (eg, delivery of fluoride, bleach or other whitening substances to the teeth).

  In other embodiments, the silicone mesh can be used in a barrier to regulate water penetration for drug controlled delivery. A mesh of polysiloxane fibers (eg, silicone fibers) can act as a barrier to regulate drug release. For example, if the drug delivery device has a large burst, a PDMS mesh (which is relatively hydrophobic) can be placed around the device to prevent or delay contact of the device with water. obtain. Furthermore, since silicone is elastic, expansion of the mesh can result in changes in its porosity and pore size, which can result in increased water to cause more drug release.

  In some embodiments, the silicone mesh can be used in pressure sensitive adhesive bandages. In this embodiment, the silicone microfiber or nanofiber is electrospun from silicone having adhesive properties. The mesh can then be applied to the skin and adheres well, but provides water and air permeability to promote natural skin function and health. This can be used in a dressing as part of a wound dressing or for a drug delivery patch.

  In some embodiments, silicone mesh can be used for oil-water separation since silicone is known to be relatively hydrophobic. If there is a high pore volume fraction, the silicone microfiber or nanofiber mesh will separate the oil from the water. The silicone can be surface treated, functionalized, or doped with additives to make it more lipophilic or hydrophobic. In this application, the silicone mesh can be used as a filter or placed in an oil-water mixture to remove the oil or separate the oil from the water. This can be extended to other systems involving hydrophilic and hydrophobic materials or phases. Since the mesh is very elastic, can the mesh be stretched to clear / remove the oil from the holes for oil recovery and / or reuse of the mesh? Can be squeezed or compressed. In addition, silicone can also absorb organic solvents and can be used to separate aqueous media from organic solvents as well. The high surface area of the microfiber mesh makes it particularly efficient and interesting for these applications.

  In some embodiments, the silicone mesh can be used in textiles. Silicone microfibers or nanofibers can also be used in textile applications where high elasticity, durability, and permeability are desired. In other applications, the hydrophobic or liquid water repellent nature of the silicone microfiber or nanofiber mesh (due to the structure) provides protection from liquid while still allowing air permeability, allowing the skin to “ Can be used to allow “breathing”.

  In various embodiments, the composite fibers can be collected into conditioned fiber bundles such as threads. These yarns can be used (eg, sutures) or further processed (multiple twisted yarns that become ropes, woven sheets, tubes or other woven yarns that become other shapes, stents, scaffolds or other tubular shapes A plurality of braiding yarns that become a structure), acting as strong, elastic fibers.

(Fibers with polymer core and hydrophilic or hydrogel sheath)
Standard materials can be produced by forming various hydrophilic or hydrogel materials around various polymer core materials, which act as reinforcing materials for the hydrophilic or hydrogel materials. The encased polymeric material may impart material properties (mechanical, chemical, thermal, etc.) specific to the hydrophilic or hydrogel material that would otherwise be impossible.

  More specifically, in some embodiments, the core-forming polymer is a hydrophilic or hydrogel-forming polymer to create a novel composite fiber having a polymer fiber core that is at least partially surrounded by a hydrophilic or hydrogel sheath. Are simultaneously electrospun. The electrospinning can be accomplished by needleless electrospinning, coaxial electrospinning, slit surface electrospinning, or any other suitable technique known in the spinning art. The result is a composite fiber with mechanical and hydrating properties that is different from any material alone. These composite fibers can be collected, formed, or processed into various shapes (eg, tubes, meshes, threads, etc.) for use as medical devices or other products.

  Any suitable hydrophilic or hydrogel-forming material can be used as the sheath polymer, and as with the choice of polymer core material, the hydrophilic or hydrogel-forming material is selected to suit the particular purpose of the composite fiber. obtain. For example, for the hydrophilic or hydrogel polymer sheath, cross-linked PVP, PEO, PVA, and hydrophilic polyurethanes, among other polymers, and xerogels, airgels, etc. among many other possibilities can be used. Other hydrogel polymers include crosslinked versions of hydrophilic polymers such as those listed above.

  Similarly, any suitable polymer can be used for the core-forming polymer, depending on the mechanical or chemical needs at hand. In some embodiments, the fibrous core is formed using a relatively hydrophobic polymer. While certain embodiments use covalently crosslinked silicon-based organic polymer cores (eg, polysiloxanes such as PDMS), the core polymer is covalently crosslinked to act as reinforcing fibers. There is no need to be done. Thus, in other embodiments, thermoplastic polymers (eg, polyurethane, PLGA, PCL, nylon, polystyrene, acrylic polymer, polypropylene, polyethylene, and fluoropolymers, among others) can be used as core reinforcing fibers.

  Polyurethanes represent a broad class of polymers with a wide range of properties and can therefore serve as core and / or sheath materials in connection with the present disclosure. For example, the thermoplastic polyurethane core can be at least partially surrounded in a hydrophilic or hydrogel polyurethane sheath. Many polyurethane materials exhibit physical crosslinking and do not require a separate crosslinking step. Such materials can be used, for example, in conjunction with, among other things, melt-based or solvent-based spinning processes.

  We relate this to the chemical nature of the polyurethane by co-electrospinning a hydrophilic polyurethane sheath around a more hydrophobic polyurethane core as detailed in Example 4 below. The concept was shown. The resulting composite fibers have different mechanical and hydration properties from either material alone.

  More specifically, a composite made of a mechanically strong polyurethane core and a hydrophilic polyurethane sheath was created. The particular technique used was slit surface core-sheath electrospinning. As previously shown, electrospinning creates fibers with small diameters (micrometers or nanometers) that provide additional benefits and functionality (eg, softness, high surface area, adaptability). However, suitable fibers can also be produced using other techniques, including melt spinning, melt electrospinning, and centrifugal fiber spinning, among other fiber forming techniques.

  As noted above, PDMS prepolymers are difficult to electrospin due to their low molecular weight and fluidity that do not allow sufficient polymer chain entanglement to form fibers. Furthermore, because the silicone prepolymer remains fluid until crosslinked, good fiber formation may not occur if the fibers are spun without any method of preserving the fiber structure. The inventors have overcome this difficulty by coaxial electrospinning to surround the PDMS prepolymer and crosslinker within the polymer sheath, especially for micro-sized and nano-sized fibers. In certain embodiments, hydrogel polymers are used as the polymer sheath material. For example, in Example 3 below, the core polymer is cross-linked PDMS and the polymer sheath is cross-linked polyvinyl pyrrolidone (PVP).

  In some embodiments, the crosslinking of the hydrogel-forming polymer is modified to suit the core material as well as the desired properties of the composite fiber. In some embodiments, hydrogel crosslinking is initiated by applying heat along with core crosslinking. For example, the core polymer can be cross-linked PDMS and the polymer sheath can be cross-linked polyvinyl pyrrolidone (PVP). Both are crosslinked by the application of heat (see, eg, Example 3). In other embodiments, methods for initiating crosslinking of the hydrogel polymer (and / or core polymer) include UV or γ irradiation, freeze / thaw cycles, supercritical drying, and the like. In yet other embodiments, a physically cross-linked hydrogel is selected (see, eg, Example 4). All of these variations of hydrogel compounds are within the scope of this disclosure.

  The main benefit of this aspect of the present disclosure is that an elastic, durable, biocompatible and mechanically stable construct can be provided for the hydrogel so that many of the potential benefits of the hydrogel are greater It can be used in applications that require mechanical integrity. Another benefit is provided a method for forming core-hydrogel fibers, which is due to the distinct cross-linking properties of the polymer selected as the core-forming polymer and / or the sheath-forming polymer. No process is required.

  As previously indicated, small diameter fiber meshes provide, among other things, higher surface area, high permeability and lower pore size than meshes made from larger diameter fibers. Thus, the present disclosure provides a material that combines the benefits of hydrogel with a small diameter fiber mesh.

  As described elsewhere, these core-hydrogel fibers can be collected, formed into various shapes (eg, tubes, meshes) for use as medical devices or other products, or Can be processed.

  Other materials can also be incorporated into the core or sheath polymer to modify or obtain new properties. For example, water-absorbing particles can be included to further improve water retention capacity, or agents that elute to provide another benefit.

  Thus, in some embodiments, excipient materials are incorporated into the fibers to increase water swelling and retention capacity. Excipient materials include cross-linked hydrophilic polymers such as PVP, cellulose, gelatin and starch, among others. These materials can be incorporated when the polymer is dissolved in the sheath or core during electrospinning. Alternatively, they can be included as particles that are not soluble or only partially soluble in the solvent used to produce the fibers. In this case, the excipient material is present as particles embedded in or protruding from the surface of the final fiber.

  In some embodiments, therapeutic agents (eg, small molecule drugs, anesthetics, procoagulants, anticoagulants, antibacterial agents, biologics, RNAi, genetic material, gene vectors, vaccines or particles such as silver nanoparticles) ) Are incorporated into the fiber and released when hydrated.

  For application, in some embodiments, the composite core-hydrogel fibers can be used in the leaflets of heart valves. Replacement heart valves in some cases use synthetic materials to recreate the native leaflet. Natural valve membranes are thin, very flexible and durable. In addition to these properties, the valve membrane must be non-thrombogenic. Promoting endothelium formation is one of the best ways to provide non-thrombogenic implants. The hydrogel layer sheath and the microfiber or nanofiber structure promote endothelial cell proliferation. Upon hydration, the hydrogel layer swells and fills the pores between the core fibers, thus preventing blood from passing through the valve pores. In a preferred embodiment, the core-hydrogel fiber microfiber mesh is electrospun to a thickness of between 100 and 1000 microns. The target fiber diameter is 500 nm to 10 μm. These meshes are then cut into the appropriate shape and attached to the body, which is implanted via laparotomy or minimally invasive surgery. Alternative embodiments include: functionalizing the core polymer with a protein or antibody (eg, CD34, VEGF) to promote tissue ingrowth and re-endothelialization (particularly degradation). Electrospinning on a frame (eg, polymer fiber, metal wire, outlined conductive mesh) to help form the valve membrane and / or provide attachment to the body Electrospinning to a biocompatible fibrous structure that creates a composite implant (eg, the fiber provides additional mechanical strength or various stiffness across the valve membrane); and coating or functionalizing the fiber; Reduce thrombus formation (eg, heparin).

  In some embodiments, the composite core-hydrogel fiber can be used in a stent graft coating. For example, hydrogel fibers can be used as a coating on a stent used in left atrial appendage closure. These embodiments are similar to heart valve leaflets, but the core-hydrogel fibers are electrospun onto a tubular collector to form microfiber or nanofiber tubes. A preferred mesh thickness is between 100 and 1000 microns. The target fiber diameter is between 500 nm and 10 μm. This tube can then be attached to a stent to form a stent graft. Alternatively, the fibers can be electrospun directly onto the stent. Alternative embodiments described with respect to the heart valve concept are also applicable here. The advantage is that the composite core-hydrogel fiber promotes cell ingrowth while providing an elastic, biocompatible and durable implant.

  In some embodiments, the composite core-hydrogel fibers can be used in vascular grafts. These embodiments are similar to the stent graft design, but the tube is not attached to the stent and the preferred mesh thickness range is larger (100-5000 microns). Alternatively, these tubular meshes are blood vessels (eg, autologous vascular grafts for bypass surgery) or other tubular structures (eg, in abdominal aortic aneurysms) when the mechanical properties of the natural tissue have been reduced. Acts as a strengthening cuff for as.

  In some embodiments, the composite core-hydrogel fibers can be used in bioengineered blood vessels. These embodiments are similar to the vascular graft described above, where the core-hydrogel microfiber or nanofiber mesh can be shaped into a tube and cells can be seeded ex vivo. These cells (typically fibroblasts, smooth muscle cells and endothelial cells) can undergo various conditions (eg, pulsatile flow, constant flow, etc.) to grow tissue on the graft material. (Or no flow) under a nutrient rich environment. The core-hydrogel microfibers or nanofibers can provide advantages in promoting cell infiltration and proliferation, as well as providing typical elastic properties to blood vessels. The core-hydrogel tube can be used alone or in combination with other natural materials (eg, collagen) or synthetic materials (eg, PTFE, ePTFE, polyurethane). In other embodiments, the graft is seeded with cells and transplanted without meaningful incubation or transplanted without cell seeding. In the latter case, cells derived from the host infiltrate and settle in the graft.

  In some embodiments, the hydrogel fibers are used in medical device closure applications. These mechanically strong hydrogel fibers, and the resulting meshes, threads, tubes, etc., ideally close the interface between the medical device and the body, other medical devices or other surfaces that require closure. Adapted to use. For example, they can be used to provide a closure between an implanted heart valve and a natural annulus to prevent peri-valve regurgitation. In one embodiment, the hydrogel fibers are electrospun directly onto the outer surface of the valve stent or formed into a kama mesh, thread or tube and applied to the valve stent as part of a manufacturing process. Upon implantation, the hydrogel absorbs moisture from the blood, which results in swelling and fills the space between the implant and the annulus, thus closing around the valve and leaking Disturb. The advantage compared to other hydrogels is that convenient mechanical properties and durability result in a safer and more effective product. Other applications include: providing hydrogel microfibers or nanofibers to a blood vessel that contacts the side of a stent graft, vascular graft or other medical device to provide the graft or other medical device and vascular wall Two stent graft components (eg, an EVAR graft body and an iliac artery leg extension) that are assembled together, providing hydrogel microfibers or nanofibers to the outer or inner diameter of the stent graft Providing the fibers outside the stent graft that is to be used as part of another stent graft deployment, such as a chimney, snorkel, etc .; outside the percutaneous catheter, colostomy bag, or lead Hide on the surface Providing gel microfibers or nanofibers to close between the device and the skin and / or underlying muscle, fat or fascia; fibers outside the device designed for implantation in the gastrointestinal tract Providing and preventing food from coming into contact with segments of the digestive system; providing fibers around an endoscope or laparoscopic instrument or access tube to provide temporary closure to the patient's tissue; Prevent bleeding, gas leaks or fluid leaks. Due to their application where the devices are temporary and removed, the strong mechanical properties and slippery surface of the hydrogel aid in removal.

  In some embodiments, the hydrogel fibers can be made such that they hydrate only when strain is applied (see, eg, Example 3 below). Upon hydration, the fibrous construct increases in volume. This property can be applied when creating strain-dependent closures around stent grafts and heart valve cuffs. In some cases, when the stent-graft and heart valve cuff are deployed, they do not fully contact the vessel wall or annulus, thereby creating an open space between the stent-graft and the vessel. Leaving this in turn can lead to leakage, device failure and poor clinical outcome. The hydrogel fibers can be used as an annulus or stent wrap so that the hydrogel fiber coating does not wet during delivery, but when the stent is deployed, the fiber coating is pulled, The fibers wet and swell, filling the empty space where the stent is not in consistent contact with the surrounding tissue.

  In some embodiments, the hydrogel fibers are used in non-medical sealing. For example, the core-hydrogel fibers are useful in providing a seal in non-medical applications in an aqueous or non-aqueous environment. For example, in an aqueous environment, the fibers placed between two surfaces to be sealed hydrate when in contact with water and then swell to seal the surface and flow through the microstructure. Disturb. In non-aqueous applications (eg, oil transport), the mesh hydrates upon installation, creating a seal from swelling between the two surfaces, and also leaking because it is immiscible with non-aqueous liquids. To prevent.

  In some embodiments, the composite core-hydrogel fibers can be used in arteriovenous grafts or shunts. These grafts are used in hemodialysis patients to provide better needle access for repeated dialysis. The core-hydrogel microfiber or nanofiber mesh provides a strong set of mechanical properties and also promotes cell ingrowth. Elasticity, durability, biocompatibility potential and low thrombogenicity improve the performance of these grafts. In one embodiment, the core-hydrogel microfiber or nanofiber mesh is shaped and implanted into a tube. The tube can be pretreated by functionalization or coating with other materials (eg, heparin, collagen, gelatin, growth factors) to improve integrity and cell ingrowth. In other embodiments, the core-hydrogel mesh can be combined with other natural or synthetic materials as a sheet or mesh to form a composite layered structure. This layered structure may improve the mechanical properties, inclusion of blood immediately after implantation, or long-term durability or performance.

  In some embodiments, the composite core-hydrogel fiber can be used in a hernia mesh. Core-hydrogel fibers (eg, a silicone or polyurethane core with a hydrogel sheath) can be electrospun into a flat mesh configuration having a thickness of 500-5000 microns. In order to improve the mechanical properties, composites can be formed with these fibers by electrospinning directly onto biocompatible polymer fibers in the desired configuration. These fibers can also be provided in a configuration that improves the suture performance of the mesh. In an alternative embodiment, the mesh can be functionalized to improve tissue ingrowth or integration.

  In some embodiments, the composite core-hydrogel fiber can be used in a dual coating. In neurosurgical procedures where the dura is injured, it is desirable to provide a coating to reclose the membrane. Core-hydrogel microfibers or nanofiber meshes (optionally combined with polymer membranes such as silicone, PLGA, collagen) can be used for this purpose.

  In some embodiments, the composite core-hydrogel fiber can be used in a wound dressing. The disadvantages of wound dressing include adhesion to the wound, management of wound exudate, and permeability to air and moisture (wound exudate). For example, the polymer surrounding the hydrogel (eg, silicone or polyurethane) can be electrospun into a mesh between 100 and 5000 microns. A preferred fiber diameter is between 500 nm and 10 microns. The advantage of the reinforced hydrogel is that it provides moisture to the wound environment and at the same time forms a protective layer that does not adhere to the wound. In one embodiment, the hydrogel-polymer bandage is supplied separately and the medical staff places additional gauze or other bandages in layers over the core-hydrogel fiber bandage. In another embodiment, the core-hydrogel mesh is combined with gauze or other backing material as part of the final product to help absorb fluid and protect the wound. In still other embodiments, the core-hydrogel mesh can be fabricated with therapeutic agents (eg, antibiotics, antifungal agents, topical analgesics, antiseptics (eg, iodotin), etc.). In yet another embodiment, the hydrogel-polymer mesh is fabricated for use in negative pressure closure therapy. In this case, the mesh is sized to be compatible with these devices and placed in the wound environment as negative pressure is applied. This high permeability and porosity allows exudate removal and non-adhesive bandages when the bandage must be removed. The hydrogel sheath or core polymer may also be useful in controlling the release of therapeutic agents to the wound (eg, antibacterial agents, antibiotics, silver ions, growth factors, analgesics, anesthetics, debridement compounds Or an enzyme).

  In some embodiments, the composite core-hydrogel fiber can be used in hemostatic applications. For hemostatic applications, the device is configured much like a wound dressing, but the hydrogel-polymer microfiber or nanofiber is fabricated with a thrombus promoter (eg thrombin, kaolin, chitosan, fibrin). Or surface modification. The fibers provide a non-adhesive bandage that can be easily removed. In addition, its high permeability and porosity allows blood to penetrate and a significant amount of surface area to come into contact with the thrombus formation promoter. This open structure also allows clotting factors to diffuse back into the wound and promote clot formation. This material can also be integrated as a non-adhesive layer on other bandages (eg, combat gauze); for this application, the fibers may or may not be manufactured with a thrombus formation promoter. It does not have to be.

  In some embodiments, the composite core-hydrogel fiber can be used in filtration. The composite core-hydrogel fiber mesh can be used as a filter for air, gas, liquid, slurry or particles or as part of a filter. In particular, its low pore size and high permeability of electrospun microfiber or nanofiber mesh is desirable for filters. Furthermore, if the fiber is elastic, its elastic nature provides a way to clean the filter. Simply stretching the material biaxially increases its pore size from the surroundings or otherwise. A gas or liquid back flow then provides a way to clean the pores from debris or other materials. In a similar manner, a cake formed on the filter's uptake side can be easily removed by extending the silicone mesh and dropping the cake. The core-hydrogel microfiber or nanofiber mesh is useful as a filter, barrier or separation membrane to distribute the oil content in water because it can be characterized by high strength and hydrophilicity. The core-hydrogel microfiber or nanofiber mesh can be used alone (preferred thickness is 100 microns to 1 cm). Alternatively, the core-hydrogel mesh can be constructed as part of a layered filter using other commercially available filter materials. In this case, the core-hydrogel can be electrospun directly onto another material, placed on the other material during assembly, or a large opening to provide wire or mechanical support Can be electrospun onto other fiber meshes having

  In some embodiments, the composite core-hydrogel fiber can be used in drug delivery. The hydrated core-hydrogel composite is substantially non-porous, but can act as a conformal layer. In one embodiment, the core-hydrogel material is inserted into the target delivery region and then expanded with a gas or other fluid (eg, a drug-containing solution) to accommodate the internal structure of the target region. Adaptive contact between the hydrogel and the surface results in efficient drug delivery. Alternatively, when a specific expansion limit is reached upon expansion, the pores are extended and allow the drug solution to be released. Once contracted, the pores are closed back so that inhibition of drug delivery to the area is not targeted during removal of the device. This approach is particularly applicable for delivery of therapeutic agents into cavities and lumens such as sinusoidal spaces.

  In various embodiments, a drug can be incorporated into the core-hydrogel microfiber or nanofiber for delivery to a patient. For example, a core-hydrogel fiber mesh can be formed with a drug in a core-forming solution and placed on the skin for dermal or transdermal delivery. The fiber meshes of the present disclosure provide a means of targeted delivery to difficult openings such as sinus cavities, intestinal walls or ear canal due to the ability to open the balloon for adaptive delivery . It may be implanted alone or held in place using another medical device such as a stent.

  In some embodiments, the composite core-hydrogel fibers can be collected into conditioned fiber bundles such as yarns. These yarns can be used (eg, sutures) or further processed (multiple twisted yarns that become ropes, woven sheets, tubes or other woven yarns that become other shapes, stents, scaffolds or other tubular shapes A plurality of braided yarns that become structures), acting as strong, elastic hydrogel fibers. These configurations can be deployed into new medical devices (eg, hydrogel catheters, introducer sheaths, guidewires, vascular grafts, hernia meshes, etc.).

  In some embodiments, the composite core-hydrogel fiber can be used in textiles. Core-hydrogel microfibers or nanofibers can also be used in textile applications where high elasticity, durability, water absorption and permeability are desired.

  In some embodiments, the composite core-hydrogel fiber can be used in tissue engineering applications. Hydrogels allow free diffusion of oxygen, nutrients, etc., desirable for these purposes. This property is further enhanced because diffusion can occur not only throughout the volume of the hydrogel but also through the porosity created by the fiber network. Hydrogels are widely used in tissue engineering applications due to their promising biocompatibility and hydration properties. The main benefit of the present disclosure is that fibrous hydrogels allow better cell binding and integration to form 3D scaffolds. The hydrogel sheath allows for cell attachment and inward growth, which is ultimately broken down, while the core polymer fiber provides more permanent mechanical support. A specific example of this application is hyaline cartilage repair, where the hydrogel sheath provides a biocompatible scaffold for stem cells to bind and differentiate into chondrocytes while its porosity is collagen and Provides space for ECM component chondrocyte secretion.

  In some embodiments, the hydrogel fiber is 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 natural tissue for a more natural appearance or feel. The fibrous nature integrates with the surrounding tissue so that the swelling agent stays in place and does not slip. Furthermore, the hydrogel can be made to be non-bioabsorbable and thus maintain its ability to expand over time.

  In some embodiments, the composite core-hydrogel fiber is used as a medical electrode. The swelling properties of hydrogels allow for adaptive and intimate contact with tissue that can reduce electrical impedance and improve electrode performance. Furthermore, to improve conductivity, the core material can be composed of a conductive polymer or can include conductive particles or ions.

  The balloon formation and hydration ability of the composite core-hydrogel fiber is a unique property that can be used prior to excising tissue through the use of microwaves. For example, for ablation in body cavities (eg, endometrium, left atrial appendage) or for excision to irregular surfaces (eg, liver, esophagus, sinuses), composite core-hydrogel fiber meshes Inflated with a gas (eg, carbon dioxide) and adapted to contact the tissue. The application of microwaves from a source in the balloon heats the water in the hydrogel membrane that is in close and adaptive contact with the cavity or tissue surface to thermally ablate the surrounding tissue. To do.

  This same technique can be used with other ablation approaches (eg, inflating the balloon with hot water (eg, inflating the balloon with hot water or other heated liquid), chemical (eg, an ablation agent in hydrogel fibers) or freezing. Can be extended to cyroblation (eg, approaches with a cooling source or liquid nitrogen used to cool the balloon).

  The composite core-hydrogel fibers of the present disclosure can also be used to embolize body cavities. The composite structure provides a fiber or coil that can be inserted into a patient using techniques known to those skilled in the art. The properties of the hydrogel then cause the fibers to swell and completely fill the body cavity or aneurysm lumen. There are two important advantages here: 1) the combination of fiber strength and high swelling ratio, and 2) the ability to form very small fibers or coils and / or flexible implants.

Example 1. Fiber with PDMS core and PLGA sheath
Core / sheath fibers are fabricated according to the use of a high throughput core-sheath needleless electrospinning facility. The sheath polymer system was 3.5% by weight 85/15 poly (L-lactic acid-co-glycolic acid) (PLGA) in 6: 1 (by volume) chloroform: methanol solvent. The core polymer consisted of PDMS (a two component liquid system consisting of part A (prepolymer) and part B (crosslinker), commercially available from Sylgard 184, Dow Corning) mixed at a 10: 1 mass ratio. While the sheath solution flow rate was set at 200 ml / h, the core flow rate was set at 20 ml / h. The fibers were deposited on a ground collection plate and collected from there. The fabricated mesh was then placed in an oven at 100 ° C. (to promote curing) for 3 hours and then immersed in chloroform for 1 hour to dissolve the PLGA sheath. The PDMS fiber mesh swelled to some extent upon exposure to solvent, but contracted back to its original size after solvent evaporation. FIG. 1 shows a cross-sectional image of the PLGA / PDMS sheath / core fiber after curing. Various polymers in the sheath / core configuration can be observed. FIG. 2 shows PDMS fibers after removal of the sheath layer. PDMS fibers were made to be between about 1-5 microns in diameter. However, as described elsewhere herein, the diameter of the core PDMS can be adjusted by adjusting the electrospinning parameters.

Example 2. Additional fibers with PDMS core and PLGA sheath
Core-sheath fibers are electrospun with 50/50 poly (D, L-lactic acid-co-glycolic acid) (5050 PDLGA) as a sheath covering a PDMS (Sylgard 184) core as described in Example 1. did. The sheath solution was 11 wt% 5050 PDLGA in hexafluoroisopropanol (HFIP). While the flow rate of the sheath solution was set to 10 ml / h, the core solution flow rate was set to 1 ml / h. The fibers were subsequently placed in an oven at 60 ° C. for 24 hours to cure the PDMS in the fibers. FIG. 3A shows the core-sheath structure formed (in cross section). The fiber diameter was measured with respect to top and cross-sectional views. The total fiber diameter of the fiber was about 7 microns (see FIGS. 3A and 3C), while the core PDMS diameter was about 4.5 microns (see FIGS. 3B and 3D). about). The 5050 PDLGA sheath was degraded by immersing the mesh for 7 days at 37 ° C. in a 12 pH buffer consisting of 1.5% sodium phosphate, 0.1% boric acid, and 0.08% citric acid. Removed under accelerated conditions. 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 provide different properties than those formed from the classic method of constructing PDMS as a cast film. The contact angle of the electrospun PDMS-only mesh was measured to be 110 °, while the PDMS cast film had a contact angle of 104 °. FIG. 4 illustrates the hydrophobic and lipophilic nature of a PDMS mesh formed using the electrospinning process of the present disclosure. Water drops (left) placed on the mesh remain beaded, while oil drops (right) wet the mesh and can move through the pores of the mesh. FIG. 5 shows the mechanical properties of the mesh compared to cast PDMS film. The data shows that the PDMS fiber mesh exhibits significantly different mechanical properties than the cast film. Compared to the cast PDMS film, the modulus of the mesh is significantly lower (0.2 MPa vs. 2.0 MPa), while its maximum load expansion is significantly higher (300% vs. 122%).

  6A-6C show cross-sectional micrographs of electrospun fibers of the present disclosure having PDMS in the core and 5050 PDLGA in the sheath. The electrospinning process, as shown in the cross-sectional views of FIGS. 6A-6C, produces sheath: core flow rates 10: 1, 10: 0.25, and 20: Performed at 0.25 ml / h.

Example 3. Fiber with PDMS core and PVP sheath
The core / sheath fibers were mixed with a sheath polymer solution of 8 wt% PVP (polyvinylpyrrolidone) in TFE (trifluoroethanol), while Sylgard 184 (part A (prepolymer) and B) mixed at a 10: 1 mass ratio. This was prepared using a core polymer solution consisting of two parts (a two-component liquid system consisting of a crosslinking agent). The sheath flow rate was set at 10 ml / h while the core flow rate was set at 2 ml / h.

  The mesh was collected on a PTFE coated aluminum shim and then cured at either 100 ° C. or 150 ° C. for 24 hours. The mesh was then removed from the aluminum shim and any uncrosslinked PVP was immersed in deionized water that was not dissolved in water. The remaining PVP was crosslinked as a strong sheath around the silicone fiber core. This then formed a hydrogel and swelled to> 200% of its initial mass (the amount of swelling is proportional to the degree of crosslinking of the PVP (and hence the curing temperature)). In particular, for the 100 ° C. sample, the swelling (by mass) was measured at 242% ± 48%, whereas for the 150 ° C. sample, the swelling (by mass) was measured at 401% ± 76%. It was.

  Gel fraction data (% hydrogel) was generated. For the 100 ° C. sample, the gel fraction was measured at 64% ± 1%, while for the 150 ° C. sample, the gelation segment was measured at 98% ± 3%. These data show that during the water extraction of uncrosslinked PVP, the 100 ° C. cured sample lost about 40% of its mass while the 150 ° C. sample maintained almost 100% of its mass. Indicates. This suggests that the PVP sheath can be almost completely crosslinked at 150 ° C. and only partially crosslinked at 100 ° C.

  Similar conclusions can be drawn by cross-sectional analysis with SEM, as illustrated in FIG. 7, which shows the following: (A) The core consists of fully cured PDMS and the sheath is PVP cured at 100 ° C. SEM cross-section of the core-sheath fiber; (B) SEM cross-section of the same fiber as (A), except after undergoing water extraction and removal of non-crosslinked PVP; (C) Core fully cured SEM cross-section of a core-sheath fiber consisting of a modified PDMS and the sheath being PVP cured at 150 ° C .; (D) except after undergoing water extraction to remove uncrosslinked PVP (C) SEM cross section of the same fiber. Prior to hydration, the two cure temperature samples show the same in core fiber diameter (about 6 μm) and sheath thickness (about 1 μm). However, after hydration and subsequent drying, the sheath appears to have been almost completely removed in the 100 ° C. sample while remaining intact in the 150 ° C. sample.

Analysis by FTIR can indicate the presence of PVP in the sample by the presence of an amine peak around 1650 cm ⁻¹ . Furthermore, since PDMS does not absorb water, the presence of a broad peak around 3400 cm ⁻¹ indicates O—H bonding and thus water absorption by the sample. This only occurs when cross-linked PVP is present. FIGS. 8 and 9 show the spectra obtained from silicone fiber hydrogels in wet and dry states, respectively, when compared to pure PDMS and pure PVP cured at temperatures of 100 ° C. and 150 ° C., respectively.

  In a comparison of the spectrum of PVP-PDMS cured at 100 ° C. with that of pure PDMS and pure PVP, it can be seen that little PVP remains in the sample after the first water extraction. Little remaining PVP can only be detected when the sample is in a hydrated state. Dry PVP-PDMS hydrogels are almost completely comparable to pure PDMS and therefore essentially do not absorb moisture from the atmosphere. This confirms the observation from the mass loss data and SEM that only essentially undetectable amount of PVP remains in the sample, but it still behaves as a hydrogel.

  In contrast to the spectrum of PVP-PDMS hydrogel cured at 100 ° C., the spectrum of the sample cured at 150 ° C. shows an amine peak and absorbed water even when dry. This is further evidence that almost all of the PVP is cross-linked at this high temperature and remains in the sample after the initial water extraction.

  The mechanical properties of the fiber hydrogel are dramatically enhanced due to the presence of the silicone microfiber structure. The tensile properties of hydrogels are rarely reported, but are difficult to find due to their poor mechanical stability. With silicone fiber reinforcement, the arrayed gastric mud gel has a larger surface area for wetting while maintaining mechanical integrity and strength. Furthermore, the presence of a cross-linked hydrogel layer on the silicone provides another layer to support the physiological corn fibers and increases the overall strength of the composite. FIG. 10 shows a comparison of various formulations of hydrogel-silicone microfiber composites, both wet and dry. The sample was cut into 20 mm squares and tested on an Instron® system at a speed of 50 mm / min. It can be seen from this data that the temperature of cross-linking can affect the tensile strength and modulus of the material.

  This data also supports the earlier stated conclusion that samples cured at 100 ° C. have little cross-linked PVP, given that the mechanical properties do not appear to be affected by the hydration state of the material. Conversely, PVP-PDMS cured at 150 ° C behaves strongly differently when dry than when hydrated. In the dry state, the material has a coefficient approximately 200 times greater than that in which it is hydrated (ie, 75 MPa vs. 0.4 MPa). It also has a much shorter elongation at break (7% vs. 140%). When the material is in a hydrated state, the PVP swells and its mechanical properties are greatly caused by PDMS fibers.

  Another characteristic of silicone core / hydrogel sheath fibers is that their mechanical characteristics can vary depending on whether the fibers are wet or dry. For example, dry mesh has high air permeability, high porosity and is opaque. Conversely, a hydrated mesh has low air permeability (because the swollen hydrogel fills the pores), high water permeability, and is optically transparent. When the mesh reaches its expansion limit, the hole opens and gas or liquid flows through the mesh, as opposed to when it bursts. This expandability is also affected by the curing temperature. This is because the fiber elongation depends on the curing temperature.

  As shown in FIG. 11A, a “balloon” was formed from a mesh of hydrated PVP-PDMS fibers cured at 100 ° C. Assuming spherical expansion, the volume expansion rate is approximately 800% when the balloon reaches maximum expansion (see FIG. 11B). At maximum expansion, the balloon does not rupture, but only becomes air permeable and air leaks out of the hole. The air permeability and porosity of the hydrated mesh can be increased when the mesh is extended and opened. Due to the reduced permeability of the hydrated hydrogel mesh, the material can retain air or water and expands to a very high volume while still maintaining mechanical integrity. In addition, the microfiber provides a flexible balloon that can conform to irregular surfaces, cavities or containers. Compared to a pure PDMS fiber mesh (which only expands about 100% before rupture), the effect of very low amount of cross-linked PVP on the surface is very significant. The PVP-PDMS hydrogel cured at 150 ° C. expands to 450% before becoming air permeable, indicating that the curing temperature and thus the amount of crosslinked PVP in the sample provides this property.

Due to the peculiarity of pairing a hydrophilic hydrogel sheath and a lipophilic core fiber, the PVP-PDMS fiber mesh swells both in water and in oil. Table 1 shows a comparison of the swelling properties of the PVP-PDMS mesh in DI water and vacuum pump oil at various cure temperatures. A mesh cured at 100 ° C. absorbs approximately the same amount of oil as the mesh absorbs water due to the presence of PDMS fibers and the presence of a very small amount of PVP on the surface. On the other hand, a mesh cured at 150 ° C. absorbs much more water than it absorbs oil. This is because much more cross-linked PVP is present in this sample.

  In addition to the differences in swelling, balloon formation and mechanical strength between samples cured at 100 ° C. and samples cured at 150 ° C., these samples also show that after their initial hydration (extraction) and drying There is a clear difference in wettability. The PVP-PDMS fibers cured at 150 ° C. absorb water rapidly and hydrate completely without any further handling or manipulation. On the other hand, meshes cured at 100 ° C. are more hydrophobic in their dry, unstretched state. They are manipulated (eg, stretched) to hydrate the mesh. In this regard, when water is first applied to the dried mesh, the water forms beads on the surface. However, as the sample is extended and manipulated, it is fully hydrated.

Example 4. Fibers with polyurethane core and hydrophilic polyurethane sheath
In this example, slit-surface-core-sheath electrospinning is used, where a hydrophilic aliphatic polyether-based thermoplastic polyurethane (HLPU) is used as the sheath material while being more mechanically stronger, more A hydrophobic aliphatic polyether-based thermoplastic polyurethane material (HBPU) was used as the core material. The electrospinning solution was as follows: 4 wt% HLPU in TFE and 6 wt% HBPU in HFIP. Electrospinning was performed at various sheath: core flow rate ratios. At the selected flow rate ratio, the resulting fibers were composed of the following HLPU: HBPU weight ratios of HLPU and HBPU: (A) 93: 7, (B) 82:18, (C) 60: respectively. 40, and (D) 38:62.

  FIG. 12 shows SEM of fibers of each composition; the fiber diameter for all formulations was approximately 2 microns.

The mesh characterization included dimensions and hydration measurements summarized in Tables 2 and 3 below. Mechanical characterization was determined by cutting the mesh into dog bone shape and performing a tensile test using Instron® at a pulling speed of 50 mm / min. Swelling was characterized by immersing the sample in phosphate buffered saline (PBS) for at least 20 minutes and weighed after the PBS drop. Swelling was calculated as (wet weight-dry weight) / dry weight. PBS retention was determined by placing the hydrated material on filter paper and applying a weight equal to 40 mm Hg for 30 seconds. The sample was then reweighed to determine the amount of moisture lost during the test. The wet tensile strength of the various polyurethane samples is shown in Table 2 and shows an increase in mechanical properties as the amount of HBPU in the fiber is increased. Accordingly, the tensile strength can be adjusted by changing the material composition of the core versus sheath.

Table 3 shows the hydration characteristics of the various formulations, showing that sample shrinkage and swelling upon hydration was most affected by the chemical composition of the fibers. However, PBS retention did not appear to be significantly affected.

  A comparison of mechanical and hydration properties as a function of HLPU content is shown in FIGS. FIG. 13 shows that the tensile strength increases as the amount of HLPU decreases (and thus the amount of HBPU increases). However, as tensile strength increases, the amount of PBS absorption decreases as a result of the less amount of hydrophilic material present.

  Comparison of swelling (or PBS absorption) and contraction data as a function of HLPU content enhances the utility of using a core-sheath fiber structure to adjust the mechanical and hydration properties. As shown in FIG. 14, the swelling capacity increases as the HLPU content increases; however, the dimensional shrinkage of the mesh (ie, area shrinkage) also increases as the HLPU content increases. Is observed. These data show the freedom of formulation of these materials and show a correlation between the performance of the hydrogel mesh and its chemical composition.

  Performance across tensile strength, shrinkage, and swelling was optimized by changing the sheath: core ratio of the polymer material. This is very advantageous for many applications, especially medical applications. For example, a hydrogel wound dressing is cut to fit the wound size when dry. These bandages improve wound healing by providing a moisture environment and absorb excess wound exudate so that it does not leak. However, excessive shrinkage can result in a bandage that improperly covers the wound after the bandage begins to absorb fluid. As shown in FIG. 15, FIG. 16 and FIG. 17, comparable water absorption (FIG. 15) compared to commercial wound dressings such as Aquacel® (ConvaTec Inc.) or Durafibre® (Smith & Nephew). 15, see Formulation A and Formulation B), providing much stronger mechanical properties (see Figure 16, all formulas) and minimal shrinkage (see Figure 17, Formulation D) And materials with shrinkage comparable to existing products (see Figure 17, Formulation B and Formulation C) have been developed.

  In various embodiments, the mesh according to the present disclosure is annealed at an elevated temperature to improve its properties. For example, an HLPU / HBPU sheath / core fiber mesh as formed herein has been found to be less porous upon annealing. In this regard, FIGS. 18A and 18B are photomicrographs before and after annealing, respectively, of a mesh formed from HLPU / HBPU sheath / core fibers as described herein. As mesh porosity is reduced, the annealing step is accomplished by a reduction in mesh volume (and hence mesh area). Unexpectedly, it has been found that such an annealing process improves moisture retention and results in mesh expansion (rather than mesh reduction). In this regard, FIG. 19 illustrates the PBS retention values of unannealed (B standard) and annealed (B annealed) HLPU / HBPU sheath / core fiber mesh, and Aquacel® and Durafiber (in accordance with the present disclosure). The retention value of a registered trademark wound dressing is shown. As can be seen from FIG. 19, an annealed mesh material was developed that provided sul PBS retention over that of Aquacel® and Durafibre® wound dressings. In this regard, FIG. 20 shows an unannealed (B standard) and annealed (B annealed) HLPU / HBPU sheath / core fiber mesh, and Aquacel® and Durafibre® according to the present disclosure. Indicates the contraction or expansion value of the wound dressing. Thus, as will be appreciated from the foregoing, the present disclosure provides that mesh absorption, retention and shrinkage / expansion can be tailored to the application at hand.

  In addition, as shown elsewhere, the resulting small fiber size also improves softness, adaptability and results in a very high surface area. High surface area improves absorption capacity, hydration kinetics and drug release capacity, among other properties. In addition, the fiber form factor allows formation / aggregation into new form factors such as yarns, ropes, tubes, meshes and the like.

(Example 5. Fiber having a polyurethane core containing silver particles and a hydrophilic polyurethane sheath)
In this example, needle core-sheath electrospinning is used, where hydrophilic aliphatic polyether-based thermoplastic polyurethane (HLPU) was used as the sheath material, while the more mechanically stronger, more hydrophobic Aliphatic polyether-based thermoplastic polyurethane material (HBPU) was used as the core material. The electrospinning solution was as follows: 4 wt% HLPU in TFE and 6 wt% HBPU in HFIP containing 30% silver particles relative to the polymer. The resulting fibers were encapsulated with silver and exhibited the core-sheath profile shown in FIG. Silver is well known for its antimicrobial properties and such a mesh can be used for the sustained release of silver for wound dressing applications. In addition to silver nanoparticles, other embodiments include incorporating other particles and / or excipients into the core material to achieve various performance matrices. For example, crosslinked cellulose or other hydrophilic polymer can be incorporated into the core to further assist the hydration properties of the resulting fiber.

  While various aspects and embodiments have been specifically described herein, modifications and variations of the present invention are covered by the above teachings and depart from the spirit and intended scope of the present invention. Without departing from the scope of the appended claims.

Claims (30)

A multicomponent fiber comprising (a) a polymer core comprising a core-forming polymer, and (b) a polymer sheath comprising a hydrophilic polymer, wherein the core-forming fiber is more hydrophobic than the hydrophilic polymer. Component fiber. The multicomponent fiber of claim 1, wherein the multicomponent fiber is formed by a core-sheath electrospinning process. The multicomponent fiber of any of claims 1-2, wherein the multicomponent fiber ranges in diameter from 0.1 to 20 microns. The multicomponent fiber according to any of claims 1 to 3, wherein the ratio of sheath volume to core volume in the multicomponent fiber ranges from 100: 1 to 1: 1. The multicomponent fiber according to any one of claims 1 to 4, wherein the hydrophilic polymer is covalently crosslinked. The hydrophilic polymer includes polyvinyl pyrrolidone, poly (acrylic acid), poly (vinyl alcohol), poly (ethylene glycol), poly (propylene glycol), poly (acrylamide), poly (methacrylate), polysaccharide, cellulose, chitosan, The multicomponent fiber according to any one of claims 1 to 5, which is selected from alginate, carrageenan, hyaluronan, gelatin and collagen. The multi-component fiber according to claim 1, wherein the hydrophilic polymer is a hydrophilic polyurethane. The multicomponent fiber of claim 7, wherein the hydrophilic polyurethane is an aliphatic polyether-based polyurethane. The multi-component fiber according to any one of claims 1 to 8, wherein the core-forming polymer is a thermoplastic polymer. 10. The multicomponent fiber according to any one of claims 9, wherein the core-forming polymer is an aliphatic polyether-based thermoplastic polyurethane. The multi-component fiber according to any one of claims 1 to 8, wherein the core-forming polymer is a crosslinked polysiloxane. The multicomponent fiber of claim 11, wherein the polysiloxane is polydimethylsiloxane. The nonwoven fabric mesh formed with the multicomponent fiber in any one of Claims 1-12. 14. The mesh of claim 13, wherein the mesh ranges in thickness from 10 to 5000 microns and the multicomponent fiber ranges from 0.1 to 20 microns in diameter. The mesh according to any one of claims 13 to 14, wherein the mesh has a wet tensile strength coefficient of at least 0.005 MPa. The mesh according to any one of claims 13 to 15, wherein the mesh has an absorptivity of at least 10% when immersed in an aqueous medium at 25 ° C for 1 hour. The mesh according to any one of claims 13 to 16, wherein the porosity of the mesh is less than 99%. The medical article containing the mesh in any one of Claims 13-17. 13. A method for forming multicomponent fibers according to any of claims 1-12, wherein the method comprises a first solution comprising the hydrophilic polymer and a second solution comprising the core-forming polymer. A method comprising the step of electrospinning the multicomponent fiber. A multicomponent fiber comprising: (a) a polymer core comprising a crosslinked polysiloxane; and (b) a polymer sheath comprising a removable sheath forming polymer. 21. The multicomponent fiber of claim 20, wherein the multicomponent fiber ranges from 0.1 to 20 microns in diameter. The multicomponent fiber according to any one of claims 20 to 21, wherein the polysiloxane is polydimethylsiloxane. The multicomponent fiber according to any one of claims 20 to 22, wherein the sheath-forming polymer is a soluble or degradable polymer. A mesh formed by the multicomponent fibers according to any one of claims 20 to 23. 25. The mesh of claim 24, wherein the mesh ranges in thickness from 10 to 5000 microns and the multi-component fibers range from 0.1 to 20 microns in diameter. A medical article comprising the mesh according to any one of claims 24 to 25. 24. A method for forming a multi-component fiber according to any of claims 20-23, wherein the method comprises a first solution comprising the removable sheath-forming polymer and a polysiloxane prepolymer and a crosslinking agent. Electrospinning the multicomponent fiber from a second solution comprising. A method of forming a silicone fiber comprising the steps of: (a) forming a composite fiber comprising a silicone core and a removable polymer sheath; and (b) removing the polymer sheath. . 30. The method of claim 28, wherein the removable polymer is a soluble or degradable polymer. 29. The method of claim 28, wherein the fibers are electrospun into a mesh form prior to removing the polymer sheath.