BIQACTIVE HYDROGEL FIBERS
Field of the Invention
[0001] The invention pertains to hydrogel fibers comprised of a carboxy-functioalized polymer (e.g. a polyacid, such as polyacrylic acid) and a hydroxy-functionalized polymer (e.g. a polysaccharide, such as dextran); articles made therewith; and uses of same. In one practice, once said fibers are formed by, e.g. electrospinning, they are crosslinked and processed to fragment same. The resultant fibers, which can be nano- and micro- fibers having diameters ranging from hundreds of nanometers to microns, can then be applied to various surfaces to create diverse topographies, including mesh-like and bristly fibrous coatings. The articles thus formed can be used inter alia for medical applications where fine control over cell adhesion properties is desired, such as in vascular grafts, tissue engineering, implants, wound dressings and the like.
Background of the Invention
[0002 } Hydrogel fibers of small diameter are generally known, often constituted of polymeric material and made by electrospinning techniques. Articles having utility in medical applications can be comprised of such fibers when the materials of construction are biocompatible. Nonetheless, there remains a need for further refinement of small diameter hydrogel technology to e.g., enable different surface topographies, modulate cell adhesion characteristics, and to create other bioactive surface properties heretofore unavailable. Summary of the Invention
[0003] The invention is directed to hydrogel fibers having small diameter (nanometers to microns), preferably of short fiber length (nanometers to millimeters) and comprised of a cross-linked carboxy-functionalized polymer and a hydroxy-functionalized polymer, such as polyacrylic acid (PAA) and a polysaccharide, such as dextran, respectively. The invention is further directed to a method of making said hydrogel fibers; articles comprised of same; and the uses of said articles in various medical and biological contexts.
Brief Description of the Figures
[0004] FIG. 1 (a)-(i) are SEM images of electrospun dextran/PAA fibers of the invention. The fibers were electrospun using different polymer concentrations, resulting in smaller and larger diameter fibers.
[0005] FIG. 2 shows a fibrous surface coating of the invention. The image shown is a confocal image of fluorescent fiber "bristles" attached to a poly-L-lysine coated surface. The fibers used were, on average, a few microns in diameter.
[0006] FIG. 3(a) and 3(b) show an endothelial cell culture on dextran/PAA fibers in accordance with the invention. In the confocal images shown, bovine endothelial cells (actin cytoskeleton stained with rhodamine-phalloidin were cultured on fluorescent dextran/PAA fibers (range of 690nm to 2.7 microns, average of about 1.5 microns in diameter) for 1 week. Endothelial cells are able to spread and proliferate on fiber-coated surfaces. FIG. 3(b) is a montage of confocal slices; FIG. 3 (a) is an overlay image of all confocal slices of FIG. 3(b).
[0007] FIG. 4 is a graph showing the results of the ethanolamine passivation of fibers in accordance with the invention.
[0008 J FIG. 5 is a graph of fluorescence intensity fibers added to a surface (micrograms of fiber). In the slope between about 100 and 250 micrograms the fiber attached in a bristle configuration to the surface.
[0009] FIG. 6 is a graph of protein adsorption on fibers of the invention.
[0010] FIG. 7 is a graph of protein adsorption on fibers immobilized onto an adhesive (e.g. PLL) surfaces and non-adhesive (e.g. dextran monolayer) surface.
[0011] FIG. 8(a)-(f) show cell adhesion resistance of fibrous surface coatings with micro- scale fibers and high density or low density fiber coatings.
Detailed Description of the Invention
[0012] As subject of the invention, hydrogel micro- and nano- fibers are fabricated and used to create selectively bioactive fibrous biomaterials. Electrospinning is preferably used to generate fibers preferably comprised of polysaccharide (e.g. dextran) and polyacrylic acid (PAA) ranging from hundreds of nanometers to microns in diameter. The dry fibers are preferably cross-linked, making them stable as individual fiber hydrogels in aqueous solution. The long, continuous fibers can be further processed into short fiber "bristles" and attached as such to various surfaces to form articles useful in medical and biological settings.
[0013] In a preferred practice, the invention is to a hydrogel fiber comprised of a polysaccharide and polyacrylic acid. Preferably, the polysaccharide is dextran. The fiber can be produced by e.g. electrospinning the polysaccharide and PAA. In a preferred practice, the polysaccharide-PAA fiber is cross-linked by methods known in the art, preferably by thermal esterification techniques. In a particularly preferred practice, the hydrogel fiber consists essentially of a polysaccharide (such as dextran) and PAA.
[0014] In another practice, the hydrogel fibers of the invention can be immobilized onto various substrates, forming nano- and microtextured coatings, or bonded to each other to create slab-like fibrous, porous hydrogel membranes. The individual polymer components of the fibers are accepted as biocompatible materials; as a result, the fibers have utility in a variety of biomaterials applications. One embodiment of the invention is directed to a substrate having a surface wherein the surface comprises a plurality of electrospun hydrogel fibers having diameters of less than one micron; the hydrogel fibers are comprised of a carboxy-functionalized polymer and a hydroxy-functionalized polymer wherein the electrospun hydrogel fibers are cross-linked and fragmented. The fragmented hydrogel fibers are attached to a surface in a bristle configuration, but when the carboxy-functionalized polymer is polyacrylic acid and the hydroxy-functionalized polymer is a polysaccharide (e.g. dextran)., the fibers may be attached to the surface in a bristle, mesh, or other configuration. Preferably, the electrospun fibers are fragmented into a length less than about 3mm. More preferably, they have a diameter of between about 10 nanometers to about 10,000 nanometers; the electrospun hydrogel fibers preferably have been cross-linked by e.g. thermal esterification; and the length of the fibers is preferred as between about 100 nanometers and about 2mm.
[0015] In another practice, the hydrogel fibers of the invention can be chemically modified to exhibit minimal protein adsorption or cell adhesion properties. This property is useful in applications in which biofouling is unwanted (e.g. implants, controlled release devices, biofiltration). In this practice, the fibers can be used to control cell adhesion, providing a protective barrier from cellular infiltration into or onto the underlying material. In minimizing protein adsorption, the fibers can prevent biofouling from disrupting the function of a material / implant / device used in vivo or in vitro or ex vivo.
[0016] In a particular embodiment, the hydrogel fibers of the invention can be used in ameliorating, including preventing, the rejection of implanted material by the host body. In one practice, the surface coatings of the hydrogel fibers are applied to implanted materials to render same effectively invisible to the immune system, thereby minimizing or even circumventing detrimental immune responses and consequent rejection for implants, which include but are not limited to: stents, catheters, grafts, pacemakers, electrodes, organ transplants, bone screws and the like.
[0017] In yet another practice, the hydrogel fibers of the invention can be custom- modified with various bioactive components (e.g. peptides, proteins, polysaccharides, enzymes, etc). Combining specific bioactivity with the fibers, preferably when they are in a mesh configuration, results in a biomimetic of native extracellular matrix (ECM; or ECM-mimetic). Specific tissue environments can thus be custom-generated for various applications, including the promotion of cell type- selective interactions, differentiation of progenitor cells, immobilization of enzymes, and the induction of specific receptor-Hgand interactions. Other applications include in situ regeneration of tissue and progenitor cell recruitment into vascular grafts.
[0018] In still another practice, the hydrogel fibers of the invention can be adapted for and used in controlled release applications. In a preferred embodiment, the fibers themselves can be loaded with substances (drugs, bioactive compounds) and either injected or implanted without eliciting an inflammatory response. Release characteristics can be modified by controlling the rate of degradation, for example, by using different fiber morphologies, controlling the degree of cross-linking, or other chemical-related methods.
[0019] In still another practice, the hydrogel fibers of the invention can be used to facilitate wound healing, by e.g., being used as a bioactive wound dressing. In one such practice, the wound dressing of the invention comprises hydrogel fibers that mimic the extracellular matrix environment of dermal wounds, in combination with bioactive molecules. In one aspect, the wound dressing can release a multitude of bioactive agents in a temporally and spatially specific, event-driven manner to promote optimal tissue regeneration and repair of chronic wounds. In one embodiment, the wound dressing of the invention provides tissue-like structure in the wound and has chemical and physical properties that support the reservoir/carrier functionality for locally delivery of wound healing bioactive molecules. The wound dressing can effectively treat a wide variety of wounds, including chronic wounds such as, but not limited to, venous ulcers, arterial ulcers, diabetic ulcers, pressure ulcers, vasculitis and the like.
[0020] Other applications for the inventive hydrogel fibers take advantage of the increased surface area created by the fibrous surface coatings, as opposed to that of a flat surface. In the field of biosensors, fibrous surface coatings can significantly increase the available surface area for specific binding of the target molecule, resulting in a more sensitive sensor. Filtration is another application contemplated by the invention, either to remove components or specifically isolate certain molecules (similar to HPLC or dialysis device), with increased surface area increasing contact area either static or under flow.
[0021] Specific preferred practices will now be described. Fiber Composition:
[0022] A first aspect of the invention contemplates a hydrogel fiber of small diameter, e.g. less than 10 microns; preferably between about 10 nanometers (ran) and about 100,000 run; more preferably between about 20 ran to about 10,000 nm. The fiber is comprised of at least one carboxy-functionalized polymer and at least one hydroxy- functionalized polymer. Carboxy-functionalized polymers and molecules serviceable in the invention include, without limitation, polyacrylic acid (PAA), carboxy-methyl dextran (CM-dex), poly(methacrylic acid), polysaccharides such as hyaluronic acid, multi-carboxylic acid molecules such as citric acid, succinate, and malate, peptides and proteins with multiple carboxy-functionality and other like polymers having carboxylic acid (-COOH) functionality . Hydroxy-fiinctionalized polymers are those having -OH groups therein, including without limitation, polysaccharides (such as dextran), polyvinyl alcohol (PVA) and other like polymers. As appreciated by the artisan, while the number of acid groups and hydroxy groups can vary, it is preferred if the ratio therebetween is sufficient to permit cross-linking to occur to a degree wherein individual fibers are thus rendered stable to aqueous solutions.
In a preferred practice, the hydroxy-functionalized polymer is dextran of Mw of between 5K to 2MM Daltøns, preferably about 7OK Daltons; and the carboxy-functionalized polymer is polyacrylic acid of Mw between about 5K to about 2MM Daltons, preferably about 9OK Daltons. The ratio of dextran:PAA can be varied between about 0:10 and about 10:0, with a ratio of about 10:4 being preferable. As appreciated by the artisan, higher or lower molecular weight formulations of these polymers can also be used within the limits of the change in morphology of electrospun fibers that result; i.e. fibers will be either thicker or thinner in diameter or may include an increased or decreased number of round "beads" interspersed with the fibers. As will be appreciated by the artisan, lower molecular weight polymers result in lower viscosity electrospinning solutions; and that a correlation exists between electrospinning solution viscosity and fiber diameter / bead defects, whereby lower viscosity equates to smaller diameter and more bead defects. In one embodiment, the dextran/P AA fibers can be made fluorescent by substituting about 0.5wt% of dextran with about 70 Kilodalton (K; g/mol) of fluorescein isothiocyanate dextran (FITC-dextran) prior to electrospinning. [0023 j In another aspect, the invention contemplates a method of making hydrogel fibers comprising: (a) electrospinning an admixture of a carboxy-functionalized polymer and a hydroxy-functionalized polymer, as herein defined, under conditions effective to form fibers from same, said fibers having diameters less than about 10 microns; (b) optionally cross-linking the fibers of step (a), by e.g. thermal esterification, to render them stable in aqueous solution; and (c) fragmenting the cross-linked fibers step (b), by e.g. mechanical shearing, including sonication, into lengths less than 3mm, preferably less than 2 mm, more preferably between about 100 nm and 1 mm. The resulting fibers can be used to form articles of manufacture by, e.g. step (d) attaching said fibers (in a flat configuration, bristle configuration, or both), onto a surface (e.g. a positively charged surface or other suitable substrate) to form an article of manufacture or a component of same.
[0024] In another aspect, the invention relates to an article of manufacture comprising a plurality of electrospun hydrogel fibers having diameters of less than one micron, said hydrogel fibers comprised of a carboxy-functionalized polymer and a hydroxy- functionalized polymer, said electrospun hydrogel fibers being cross-linked and fragmented into a length less than about 3mm, said fragmented hydrogel fibers being attached to a surface in a bristle configuration. Preferably, said electrospun hydrogel fibers have a diameter of between about 2 nanometers to about 10,000 nanometers; said carboxy-functionalized polymer is polyacrylic acid, said hyroxy-functionalized polymer is dextran; and said electrospun hydrogel fibers have been cross-linked by thermal esterification; the length of said fibers preferably between about 100 nanometers and about 2mm. A preferred use setting for this article is a medical device. In a further aspect of the invention relates to a substrate having a surface, said surface comprising a plurality of hydrogel fibers attached thereto. In one embodiment, said fibers are comprised of a carboxy-functionalized polymer and a hydroxy-functionalized polymer. In another, said carboxy-functionalized polymer is polyacrylic acid or carboxy-methyl dextran; said hydroxy-functionalized polymer is dextran or polyvinyl alcohol; and said fibers are cross-linked, preferably by thermal esterification. The surface of the substrate can comprise the hydrogel fibers in various configuration, with bristle and mesh configurations preferred, provided preferably that when the fibers are comprised of PAA and a polysaccharide (e.g. dextran), the configuration can be bristle or mesh; and when other carboxy-functionalized polymers and hydroxy-functionalized polymers comprise the fiber, the configuration is preferably bristle. In one practice of the invention entails the proviso that when said carboxy-functionalized polymer is polyacrylic acid and said hydroxy-functionalized polymer is a polysaccharide, said fibers may be attached in a bristle, mesh, or other configuration.
[0025] One practice of the invention relates to a device for biological or medical use comprising the substrate aforesaid, e.g. said device can comprise a filter, a membrane, a tissue graft, wound dressing, or an implant, e.g. for controlled release of drug substances. The device can also comprise fibers that have been modified to exhibit decreased cell adhesion, or that have been modified with one or more peptides, proteins, polysaccharides or enzymes. When the device comprises a wound dressing it is preferred if the bioactive agents are also utilized with same, such as growth factors, cytokines, bioactive peptides, or combinations thereof. Preferably, the wound dressing is formulated into a moisturizing dressing, an absorptive dressing or a wound filler. The wound dressing contemplated by the invention can be used to treat, without limitation, venous ulcers, arterial ulcers, diabetic ulcers, pressure ulcers, vasculitis and the like. Implants contemplated by the invention include without limitation, stents, catheters, pacemakers, electrodes, bone screws and the like.
[0026] Other aspects and embodiments of the invention are described hereinbelow using the preferred PAA/dextran fiber. Notwithstanding, it is understood that this is for convenience only, and that these descriptions are applicable to other compositions and practices, all of which are within the ambit of the invention.
Fiber Generation:
[0027J Fibers can be generated to the diameters specified by methods known in the art, all of which are contemplated by the invention. In a preferred practice, the fibers are electrospun using commercially available technology and techniques as known to the artisan. See e.g. Zong, XH; Kim, K; Fang, DF; Ran, SF; Hsiao, BS; Chu, B. 2002. Structure and Process Relationship of Electrospun Bioabsorbable Nanofiber Membranes. POLYMER 43 (16): 4403-4412, incorporated herein by reference. [0028] Without limitation, electrospinning is performed using water as solvent; or using water a the primary solvent in the electrospinning solution in concert with other solvents, such as DMSO, DMF, ethanol and the like. Water when present can optionally have dissolved therein anionic or cationic salts, such as sodium chloride, calcium chloride, potassium chloride and the like, and/or dissolved proteins or peptides, and/or dissolved small molecules, such as fluorescent dye molecules and the like. Electrospinning can be performed using polymer solution concentrations of between about 0.4 to about 1.2 grains of polymer dissolved in 1 ml of solvent. A low humidity environment is preferred for the electrospirming, e.g 0% to about 30% humidity. Voltage applied to the electrospinning solution is between about 15KV to about 30KV. Solution flow rates are between about 1 μL/min to about 50 μL/min from the capillary tip. A metal capillary of about 0.01 inch to about 0.05 inch in diameter is used to extrude the polymer solution. The electrospinning is performed with the target located about 5 cm to about 20 cm from the capillary tip.
[0029] In one preferred practice of the invention, an aqueous solution is prepared by dissolving a mixture of dextran and PAA polymers in deionized water. The concentration of total polymer can vary between about 0.6 to about 1.0 gram per 1 mL water. The ratio of dextran:PAA can also be varied between about 0:10 and about 10:0, with a ratio of about 10:4 being more preferred. Electrospinning can be performed at a flow rate of between about 7-15 μL/min, using a high-voltage power supply to apply 25kV to a metal capillary tip (0.02 inch to 0.03 inch inner diameter). The electrospun fibers that result are collected on a grounded metal target located between about 10-20 cm from the tip. Room humidity is maintained between about 14-16%. After electrospinning, the fibers are dried under vacuum overnight. As will be appreciated, fiber diameter is controlled primarily by varying initial polymer solution concentration, with lower concentrations generating smaller diameter nanofibers (tens to hundreds of nanometers) and higher concentrations generating larger diameter microfibers (e.g. microns). Intrafiber Cross-Linking and Stability:
[0030] Stability in aqueous solution, degree of swelling and mechanical stability can be controlled by intrafiber cross-linking density. To prevent dissolution in water, the dextran/PAA fibers can be cross-linked using techniques known in the art. Preferably, cross-linking is performed using a thermal dehydration reaction based on the procedure in Chen, H; Hsieh, YL. 2004. Ultrafme Hydrogel Fibers With Dual Temperature and pH Responsive Swelling Behaviors. JOURNAL OF POLYMER SCIENCE PART A- POLYMER CHEMISTRY 42 (24); 6331-6339, incorporated herein by reference. The dried fibers resulting from, e.g., electrospinning, are placed in a vacuum oven for between about 1 minute to about 1 month, preferably between about 15 minutes and about 1 week, more preferably between about 1 hour to about 2 hours, at between about 120° C and 300° C, preferably between about 140° C to about 210° C, more preferably about 18O0C. This induces thermal esterification between the hydroxyl groups (-OH) of the dextran and the carboxylic acid groups (-COOH) of poly(acrylic acid); i.e. an ester bond is formed between these groups. As understood by the artisan, the degree of cross- linking (number of ester bonds formed) is controlled by varying the 'mer' ratio of dextran to polyacrylic acid. In general, fibers with a higher degree of cross-linking are more stable when immersed in water, while those with lower degrees of cross-linking are less stable. For example, dextran/PAA fibers created using about a 10:1 mer ratio result in a low degree of cross-linking, and will initially swell when immersed in water, then dissolve within a short time. However, dextran/PAA fibers created using about a 10:4 mer ratio will maintain their fibrous structure when immersed in water, swelling and forming a network of hydrogel fibers that remain stable long-term (dissolution in water does not occur after >1 year, ~pH=7). As generally appreciated by the artisan, fiber stability is affected by mechanisms that cleave ester bonds, including hydrolysis, strong acids or bases, or esterase enzymes.
Fiber Fragmentation:
[0031] In a preferred practice of the invention, the fibers are fabricated into short fiber segments or "bristles." In a preferred practice to fragment the long electrospun fibers into pieces less than about 3mm, to more preferably ranging between about lOOnm to about 2mm millimeters in length, more preferably between about lOOnm and about lmm. In a preferred embodiment, the length is less than about 0.1 μm (105 nm). While any form of mechanical shearing can be used, such as forcing the fiber through a narrow-gauged syringe, sonication is preferred. In this practice, the dry fibers are removed from the metal collector after electrospinmng, cross-linked via thermal esterification, and immersed in deionized water. A sonicator probe is inserted into the fiber solution at a setting of 4 for about 10 seconds, which mechanically shears the long fibers into short fragments. The fragmented fibers, or "bristles," that result are stable between the temperatures of about 4° to about 37° C, at pH~7, and can be stored in deionized water (DI) (typically at 40C) until ready for use.
Fiber Surface Attachment:
[0032] The fibers of the invention can be attached to various surfaces as known in the art. In the practice of the invention the fibers fragmented, they may be attached in a single layer, e.g. having a "mesh" configuration wherein the fibers are attached to lie flat on the surface; in a "bristle" or "bristly" configuration wherein they protrude upwards from the surface; or they can be attached in multiple layers on a surface, to form a multiplayer coating of attached fiber fragments. Where the fibers are not fragmented, they may be attached to a surface in a single layer, or in multiple layers. In a preferred practice, where the fibers have a net negative charge, as e.g. retained after being electrospun, the fibers can be attached to a positively-charged surface, such as poly-L-lysine adsorbed onto tissue culture plastic or acid-etched glass. In a preferred practice, a fiber solution in water, typically from about 1 microgram to about 10 micrograms of dry fiber bristles hydrated in about 1 mL DI water, is added to a positively-charged surface, preferably poly-L-lysine-coated tissue culture plastic (PLL), and the fibers are allowed to settle overnight. After rinsing, the fibers remain attached. As will be appreciated by the artisan, lower concentrations of added fiber result in sparse surface coverage, with fibers tending to lie flat. Higher concentrations of added fiber result in dense surface coverage, with fibers tending to "end-attach" more often and protrude upwards to form a "bristle" or "end on" configuration, i.e. one end terminus of the fiber is attached to the surface. Fibers allowed to dry onto the positively-charged surface attach flat. As a result, a variety of surface topographies can be generated by varying fiber surface density and fiber diameter, and by attaching the fibers flat or upright, resulting in either a mesh-like or bristly surface, respectively. In consequence, overall surface topography and surface area is controlled by fiber attachment method, fiber selection, and fiber density. Once attached, the fibers are stable in solution (including e.g. PBS5 cell culture media, DI water, high and low pH solutions of weak acids or bases) and in the dry state for many months or longer. In addition, the attached fibers are very stable to vigorous rinsing. Mechanical shear such as scratching with a blunt probe removes adherent fibers. Fiber density can be controlled by manipulating the concentration of fibers in solution. Surface morphology can be controlled by allowing the fibers to dehydrate onto the surface, which results in a mesh morphology; or stay hydrated, which results in a bristly morphology.
[0033] In another embodiment, the fibers can be aligned, e.g. they can be attached to a surface such that they lie parallel to each other. This is accomplished by flowing a solution of fiber bristles in water over a positively charged surface. The fibers attach, then lie down and attach in the direction of flow.
Fiber Passivation.
[0034] In a preferred practice, the fibers are passivated, as may be desired for various use settings, e.g. where non-specific protein fouling and cell adhesion is disfavored or can betolerated only in limited fashion. Suitable passivation techniques are known in the art. In a preferred practice to passivate the fibers, electronegative -COOH groups are neutralized by attaching ethanolamine at these sites. In a typical procedure, a solution of 3OmM EDC, SmM NHS, and 0.1M ethanolamine in 5OmM MES buffer is added to a fiber-coated surface and allowed to react overnight, then rinsed thoroughly.
Selective Bioactive Modifications:
[0035] In another embodiment, the bioactivity of the fibers is modified. In a preferred practice in this regard, bioactive modification is performed similarly to passivation, i.e.instead of ethanolamine, a specific bioactive molecule containing an amine group (peptide, protein, enzyme, etc.) is added. For example, N-(3-dimethylaminopropyl)-N'- ethylcarbodiimide (EDC) and N-hydroxysuccinimide (NHS) can be used to NHS-activate carboxyl groups, the sites at which an amine group will covalently bind. As will be appreciated by the artisan, the degree of modification can be controlled by varying the concentration of bioactive molecule, or decreasing or increasing the number of NHS- activated carboxyl groups. Carboxyl groups can be totally eliminated by ethanolamine passivation, while additional carboxyl groups can be generated by incubating the fibers with bromoacetic acid (0.1 M to IM) and sodium hydroxide (2M NaOH) for a given time (1 minute to 24 hours). Lower molarity (e.g. 0.1M bromoacetic acid) solutions result in lower degree of fiber carboxylation (i.e. the generation of a smaller number of -COOH groups) and vice versa for higher molarity (e.g. IM bromoacetic acid).
Differential Bioactive Modifications between Fibers and Underlying Surfaces: [0036] In another embodiment, the fibers are differentially modified from each other, or from the underlying surface. Without limitation, one preferred method for differentially modifying fibers is sequentially attaching fibers in batches, while modifying each batch before the next batch is added. In another preferred method, fibers can be modified prior to attachment, and combinations of fibers with selected modifications can then be attached. In regard to differentially modifying the underlying background surface, a preferred practice is to apply a dextran monolayer. This passivates the background surface (which was previously a positively charged adsorptive/adhesive surface). As a result, fibrous surfaces can be created with combinations of (1) adhesive fiber, non- adhesive background (2) non-adhesive fiber, non-adhesive background (3) adhesive fiber, adhesive background (4) non-adhesive fiber, adhesive background (adhesive = bioactive, protein adsorptive; non-adhesive = minimal protein adsorption, passivated, not cell adhesive). Variations on these methods as appreciable to the artisan and are within the ambit of the invention. In yet another practice, the bioactive molecule can be included in the electrospinning solution (prior to electrospinning) and cross-linked into the fibers (e.g. as in making fluorescent fibers by incorporating FITC-dextran in the electrospinning solution).
Tunable fibrous surface coatings for implantable materials
[0037] In one aspect, this practice of the invention mediates the interfacial interaction between the body (cellular response/biological components) and underlying implanted material comprising fibers of the invention. Implants in this regard ameliorate, including preventing, rejection of the implant by the host body. As appreciated by the artisan, mediation for this purpose invokes three properties: chemical, topographical, and mechanical. In one practice of this aspect of the invention, these three properties are tunable and can be optimized to tailor the host body's response towards acceptance of the implanted material.
[0038] In one preferred embodiment, the fibrous surface coatings for implantable materials are constructed from nanofibers and microfibers of dextran and polyacrylic acid (PAA) polymers as herein described. They include dextran-PAA hydrogel fibers, or "bristles," which are tunable in size between about 400nm to about 4μm in diameter, and between about lμm to about lmm in length. Immobilizing these fibrous "bristles" onto the surface of an implantable material forms a coating that functions, e.g., as a camouflaging barrier or bridge. [0039] In accordance with the invention, when fibers thereof are implemented as a barrier coating on an implant, they are preferably customized to resist either or both protein adsorption and cell adhesion, which events typically instigate inflammatory responses leading to implant rejection. In one practice, the fibers can be modified to include chemical groups to which proteins minimally adhere. The experimental results demonstrate this protein-repellant nature (see Example 7).
[0040] Resistance to cell adhesion is' also achieved by optimizing fiber topography and coating density. Example 8 demonstrates the cell resistant nature for fibrous coatings applied at high surface density or constructed using micro-scale fiber diameters. Synergistically combining protein-resistant chemistry and cell-resistant topographies and coating densities can result in a surface coating that can be applied on an implant as a resilient barrier to the immune system. Camouflaged by this protein- and cell-resistant fibrous barrier, the underlying implanted material is able to effectively appear "invisible" to the immune system, evading recognition as a foreign body and ultimate rejection.
[0041] In another embodiment, fibers of the invention can be implemented as a bridge coating so that they are customized to interact with the host body's cells to promote natural incorporation of the material. In one practice of such coatings, the fibers are patterned with topographical and chemical cues derived from the native extracellular environment. This mimicry of the native environment can provide signals to the body's cells, inducing them to integrate the fiber-coated, implanted material within the body. Hvdrogel fibers in wounding dressing. [0042] In another embodiment, the hydrogel fibers of the invention can be used as a bioactive wound dressing, including for the treatment of chronic, non-healing wounds. Chronic non-healing wounds often involve progressive tissue loss and a disruption of the normal process of healing. The wound dressing of the invention can provide a provisional tissue structure that initiates accelerated wound healing.
[0043] The highly tunable mechanical properties of the hydrogel fibers of the invention, topography, bioactivity, and biodegradation rates, are suitable for wound healing inasmuch as (1) they permit a fine degree of control over ranges emulating that of native cellular and tissue environments; (2) the innate nature of the fiber material is non-fouling, providing an inert background for adding back specific bioactive functionality; and (3) the fibers can be applied as a surface coating or comprise the scaffold itself.
[0044] In one practice, the wound dressing of the invention comprises hydrogel fibers that mimic the extracellular matrix environment of dermal wounds in combination with bioactive molecules. These can include various growth factors, cytokines, and bioactive peptide fragments that are released in a temporally and spatially specific, event-driven manner to promote optimal tissue regeneration and repair of chronic wounds. The hydrogel fibers can also serve as a reservoir or carrier for bioactive molecules that facilitate tissue regeneration and wound healing. In practice, the hydrogel fibers can also add back tissue-like structure in the wound and have chemical and physical properties that support the reservoir/carrier functionality for locally delivery of wound healing bioactive molecules. Furthermore, biodegradation rates of the hydrogel fibers can be tailored to control specific time release of bioactive agents, thus creating a smart wound healing system. Preferably, the hydrogel fibers for this practice are dextran-polyacrylic acid (Dex-PAA) fibers ranging from about 200 nm to about 3000 nm diameter.
[0045] The wound dressing in accordance with the invention can be formulated into various wound care modalities, including but not limited to, 1) moisturizing dressings; 2) adsorptive dressing; and 3) wound fillers.
[0046] The succeeding examples are of various embodiments of the invention; other practices will be apparent to the artisan.
Example 1
[0047] Generation of fibers and cross-linking: an aqueous solution of dextran/PAA was electrospun at concentrations of 0.7 g/mL and 0.75 g/mL using a mer concentration ratio of dex:PAA of 10:4, then cross-linked at 180° C under vacuum for 1 hour. Samples were Au-sputtered, then imaged using an XL-30 Scanning Electron Microscope (SEM). FIG l(a) and FIG l(b) and FIG l(c) are the SEMs of the fibers resulting from the 0.7g/mL concentration taken at 100Ox, 2000x, and 10,000x, respectively. FIG l(d) and FIG l(e) are the SEMs of the fibers resulting from the 0.75g/mL concentration, both taken at 10,000x.
[0048] The same materials and conditions were used only with concentrations of 0.6g/mL, 0.65g/mL, 0.7g/mL and 0.08 g/mL. SEMs at 10,000x of the Au-sputtered fibers that resulted are shown in FIG l(f), FIG l(g), FIG l(h) and FIG l(i), respectively. Example 2
[0049] Fragmentation and Attachment of Fibers to a Surface: aqueous solutions of dextran/PAA were electrospun at varying concentrations, ranging from 0.6 to 0.9 g/mL, and using a mer concentration ratio of dex:PAA 10:4, then cross-linked at 180° C under vacuum for 1 hour. The resultant fibers were then sonicated and fragmented into segments, then immobilized onto a poly-L-lysine-coated (PLL) surface. FIG 2 shows a fluorescent confocal image wherein 0.5% FITC-dextran was mixed in a 0.8 g/mL dex/PAA solution prior to electrospinning, thermally cross linked, then attached to a PLL-coated surface; fiber concentration was about 200 micrograms/mL. As shown in the perspective confocal image of FIG 2, fibers attached in a relatively flat or sticking up ("end-on" or "bristle") morphology.
Example 3
[0050] Cytoskeleton of Endothelial Cells Cultured on Fibers: (a) Bovine endothelial cells (BECs) were cultured for 1 week on PLL surfaces coated with a high density of FITC- fibers (darker area of FIG 3 (a)) as made in accordance with Example 2. Cells were then formaldehyde-fixed and the actin cytoskeleton stained using phalloidin-rhodamine (lighter area of FIG 3 (a)). A confocal fluorescence image, FIG 3 (a) was then taken using a water-dip lens at 64x. As shown in the image of FIG 3(a), endothelial cells were able to spread and proliferate on the fiber-coated surfaces. FIG 3(b) is a montage comprised of the same fluorescence confocal slices used to generate FIG 3 (a). From left to right then top to bottom, slices are shown from the bottom of the surface to the top. As shown in the images of FIG 3(b), the cell is growing on top of and intermingling with the fibers.
Example 4
[0051] Passivation of Fiber -CQOH Groups: following electrospinning and cross- linking as in Example 1, the fibers retained negatively-charged carboxy (-COOH) functional groups derived from poly(acrylic acid). Toluidene blue, a small positively charged dye molecule, is assumed to bind each -COOH in. a 1 :1 ratio. A toluidene blue assay was developed to determine the presence of as well as elimination of the -COOH functional groups. In this example, 80 μg of fiber were immobilized onto poly-L-lysine (PLL) coated wells of a 24-well plate. For passivation, treated wells were incubated with 0.1M ethanolamine/50mM MES/30mM EDC/8mM NHS (pH=6) overnight while rocking. After rinsing thoroughly with DI water, all wells were incubated with 1 mL of toluidene blue (SxIO"4 M, pH=10) for 1.5 hours, then rinsed thoroughly with DI water (pH-10). Water was removed, then 500 μL of 50% acetic acid was added to unbind toluidene blue from -COOH. Three 100 μL samples were taken from each well and absorbance was read at 595 nm in a 96-well plate. FIG 4 is a graph of the results. As shown in FIG 4, absorbance readings showed that minimal toluidene blue was bound to fibers treated with ethanolamine versus untreated control fibers. This indicated that the carboxy groups of the fibers were successfully passivated.
Example 5
[0052J Fiber Immobilization Density: An increasing amount of FITC-labeled fibers were added for immobilization to poly-L-lysine (PLL) coated wells of a 24-well plate (2 cm2 surface area/well). For immobilization, fiber solutions were made in DI water in concentrations ranging from 5 to 360 μg per mL. 1 mL of fiber solution was added per well and incubated overnight, then rinsed 3x with water. Fluorescence was measured at ex/em wavelengths of 485/535nm. As shown in FIG 5, the amount of immobilized fiber increases linearly from zero to approximately lOOμg, and transitioning to a more logarithmic trend towards 360μg. The fibers attached to the PLL surface either sticking up (end-on or bristle configuration) or relatively flat. The ratio between these morphologies differed depending on the packing density at each fiber concentration. As concentration increased, the number of "end-on" fibers increased. The total immobilization amount increased up to a saturation point at which fibers are sterically hindering other fibers from the reaching the surface.
Protein Adsorption on Fibers: Example 6
[0053] Differing amounts of fiber were immobilized into poly-L-lysine (PLL) coated wells of a 24-well plate (2 cm2 surface area/well). After rinsing each well with PBS, the buffer was replaced with 500 μL of FITC-BSA in PBS (600 μg/mL) and incubated for 2 hours. Fluorescence was measured at ex/em wavelengths of 485/535 run. Initial protein adsorption, as shown in FIG 6, was measured following 3 rinses of PBS. Protein adsorption after rinsing, as shown in FIG 6, was measured following 4 hours of rocking the sample and many rinses with PBS. Fiber-immobilization increased the amount of initial protein adsorption versus the PLL-control surface. However, after thorough rinsing, all protein adsorbed into the fiber was released such that protein levels matched that of protein adsorption on the control PLL surface.
Example 7
[0054] Surface treatments were performed on 24-well plates. Each well was incubated overnight with 0.5 mL of 500 μg/mL FITC-BSA in PBS, then rinsed 5x with PBS. To determine total levels of adsorbed BSA, fluorescence intensity readings were taken at ex/em wavelengths of 485 nm/535nm both prior to FITC-BSA incubation and after rinsing. Intensity readings were normalized such that BSA adsorption on PLL = 1. The results are shown on FIG 7. FIG 7 graphs FITC-BSA protein adsorption levels on PLL and dextran monolayer surface treatments, with and without immobilized dextran/PAA fibers. That is, fibers were immobilized onto either PLL {an adhesive) or onto dextran monolayer (a non-adhesive) background surface coating. As shown FIG 7, immobilized fibers did no increase protein absorption levels as opposed to the levels measured on the background surface. On an adhesive PLL surface, immobilized fibers actually decreased protein adsorption levels. On a generally non-adhesive surface {dextran monolayer), immobilized fibers exhibited adsorption levels comparable to that of the dextran monolayer.
[0055] As seen in FIG. 7, protein adsorption was measure on four different surfaces: 1) PLL, as a control protein-adhesive substrate; 2) Dextran-PAA fibers immobilized on PLL; 3) dextran monolayer (DM)5 as a control protein-repellant substrate; and 4) dextran- PAA fibers immobilized on DM. As shown in FIG.7, a maximum amount f protein was adsorbed on PLL, the adhesive control, whereas when this PLL substrate was coated with fibers, protein adsorption decreased. For DM, the non-adhesive control, a minimum amount of protein was adsorbed on it. When the DM surface was covered with fibers, protein adsorption remained about constant. The results establish that the fibers of the invention are protein-repellant, and capable of reducing overall protein adsorption when immobilized onto a protein-adhesive background substrate. Implants that comprise protein-resistant fibers of the invention will also resist cell adhesion inasmuch as cells need surface adherent proteins to bind to a surface. Particular cells that are repelled by protein-resistant fibers of the invention include, without limitation, immune cells involved in the inflammatory-rejection response of an implant material otherwise. Example, 8
[0056] Cell adhesion resistance by fiber topography and fiber coating density were measured on six different surfaces. Fibers utilized comprised short fiber bristles, immobilized onto a protein-adhesive PLL background substrate in a mesh-like topography. FIG. 8 (a)-(f) shows confocal fluorescence images wherein light areas indicate fibers, and lighter areas (central portion of FIG. 8 (a)-(c)) indicates cells. High density fiber coatings are depicted in FIG 8 (a), (b) and (c). Low density fiber coatings are depicted in FIG 8 (d), (e) and (f). The diameters of the fibers varied as follows: In FIG 8 (a) and (d): from 475nm to 2.15μm. In FIG 8 (b) and (e): from 690nm to 2.7μm. In FIG 8 (c) to (f): from 1.5μm to 3.8μm. As shown in FIG 8 (c), wherein micro-scale fibers were immobilized at low density, even though cells were able to adhere to some degree, they did not easily traverse over the fibers, indicating effective resistance to cell adhesion by coatings comprised of micro-scale fiber diameters. FIG 8 (d), (e) and (f) show that high density, fibrous mesh surface coatings effectively prevented cell adhesion as not cells are readily visible; these particular topographies acted as barriers to cell adhesion by effectively restricting access to the underlying substrate.
[0057] Although preferred embodiments of the present invention have been illustrated in the accompanying figures and described in the foregoing detailed description, it will be understood that the invention is not limited thereto as defined by the claims herein.