WO2013019733A2

WO2013019733A2 - Methods, compositions and kits for therapeutic treatment with wet spun microstructures

Info

Publication number: WO2013019733A2
Application number: PCT/US2012/048860
Authority: WO
Inventors: Edith Mathiowitz; LAVIN (formerly DECOTEAU), Danya; Richard Hopkins
Original assignee: Brown University
Priority date: 2011-07-29
Filing date: 2012-07-30
Publication date: 2013-02-07
Also published as: WO2013019733A4; WO2013019733A3

Abstract

Methods, compositions, systems, devices and kits are provided for preparing and using a multi-layer polymeric microstructure composition for delivering a therapeutic agent to a subject. In various embodiments, the therapeutic agent includes at least one selected from the group of: a drug, a protein, a sugar, a carbohydrate, and a nucleotide sequence. In related embodiments, the composition is a fiber, a suture, a sphere, an implant, or a scaffold.

Description

Methods, compositions and kits for therapeutic treatment with

wet spun microstructures

Related Application

This application claims the benefit of provisional application serial number 61/513,470 filed July 29, 2011 entitled, "Methods, compositions and kits for therapeutic treatment with wet spun binary phase microstructures", inventors Edith Mathiowitz, Danya Decoteau, and Richard Hopkins which is hereby incorporated herein by reference in its entirety.

Government Funding

A portion of this work was supported by the National Science Foundation Materials Research Science & Engineering Program (DMR 0520651). The government has certain rig in this invention.

Technical Field

Systems, compositions, methods and kits are provided for preparing and using multilayer polymeric microstructure system for controlled therapeutic agent delivery.

Background

The delivery of hydrophilic therapeutics from small diameter wet spun microfibers is often characterized by an initial burst due to drug trapped on the surface during encapsulation (Williamson et al. 2004 Biomaterials 25:5053-5060; Change et al. 1998 Journal of Biomedical Research Part A 84A: 230-237). Rapid drug burst from microfibers is detrimental under the circumstances of a drug having a narrow therapeutic range, thereby resulting in local drug concentrations that quickly become toxic, and little advancement has been made to slow the release of therapeutics from wet spun microfiber-based delivery systems (Schakenraad et al. 1994 Biomaterials 367: 258-260). Previous work addressed double-walled microspheres with drag localized to inner or outer core for controlled release kinetics (Rahman et al. 2003 Journal of Controlled Release 94: 163-175; Pekarek et al. 1994 Nature, 367:258-260). There is an urgent need for microencapsulation and phase separation techniques which slow the release of therapeutics from wet spun microfiber-based drug delivery systems. Summary

Compositions, methods and kits are provided for multifunctional polymeric microfibers with prolonged drug delivery and structural support capabilities.

An embodiment of the invention provides a wet spun microfiber composition having at least one polymer such that the composition includes a porous multi-layer polymeric microstructure, and further includes at least one encapsulated therapeutic agent, such that the therapeutic agent is located in an inner core of the microstructure and is controllably releasable from the composition, and such that the microfiber has a degree of crystallinity at least 10% greater than that of control polymer prior to wet spinning. In a related embodiment of the composition the polymeric microstructure is multi- layered. In other embodiments of the composition, the at least one encapsulated therapeutic agent is located in an inner core of the multi-layered microstructure.

In various embodiments the composition has a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

In related embodiments of the composition an encapsulated first therapeutic agent is dexamethasone. Various embodiments of the composition further include an encapsulated second therapeutic agent. For example, the second therapeutic agent includes at least one selected from the group of: a drug; a protein, for example, Nog (Noggin); a peptide; a sugar; a carbohydrate; and a nucleotide sequence. For example, the nucleotide sequence includes a vector. In related embodiments the protein is at least one selected-from the group of: a growth factor, an immunoglobulin, an enzyme, and a peptide antibiotic.

In various embodiments of the composition the polymers are at least one of poly-1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP).

According to an aspect of the invention, the composition comprises at least about 75% of the initial tensile strength for at least about five weeks.

Another embodiment of the invention provides a method of producing a wet spun microfiber composition having a porous multi-layer polymeric microstructure including the steps of mixing at least one polymer and at least one therapeutic agent with a solvent to form a solution; and wet spinning the material by phase inversion, thereby producing the

microstructure, such that the composition has a degree of crystallinity which is at least 10% greater than that of control polymer prior to wet spinning.

In various embodiments of the method a first therapeutic agent is dexamethasone.

Related embodiments of the method further include mixing the polymer solution with a second therapeutic agent prior to wet spinning. For example, the second therapeutic agent is selected from the group of: a protein a peptide, a sugar, a carbohydrate, a nucleotide sequence, and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anti-coagulant; an anti-tumor; an anti-viral; an anti-bacterial; an anti-mycobacterial; an anti-fungal; an antiproliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a nonsteroidal anti-inflammatory agent (NSAID).

In related embodiments of the method the solvent includes at least one of

dichloromethane and tetrahydrofuran.

In embodiments of the method, wet spinning includes loading the material into a syringe, and dispensing the material into a coagulation bath, such that the coagulation bath includes a non-solvent, thereby obtaining phase inversion. For example, the coagulation bath includes petroleum ether.

Embodiments of the method include selecting the solvent and the non-solvent having different solubility parameters, such that the difference between the solubility parameters affects the rate of solidification of the polymer, the extent of solvent induced crystallization of the polymer, and the degree of crystallinity of the composition. The difference between the solubility parameter of the solvent and the solubility parameter of the non-solvent is selected from one of the following: less than about 12 units, less than about 10 units, less than about 9 units, less than about 8 units, less about than 7 units, less than about 6 units, less than about 5 units, less than about 4 units, less than about 3 units, less than about 2 units and less than about 1 unit. In various embodiments, the difference is less than about 2-4 units, less than about 4-6 units, less than about 6-8 units, or less than about 8-10 units. In various embodiments, the difference is varied to modulate rate of crystallization of the composition. In a related embodiment, the polymers include a polymer matrix or a composite material. For example, the polymer matrix is bioabsorbable.

Another embodiment of the invention is a method of treating a subject having a medical condition including, contacting the subject with a wet spun microfiber composition having at least one polymer such that the composition includes a porous multi-layer polymeric microstructure, and further includes at least one encapsulated therapeutic agent, such that the therapeutic agent is located in an inner core of the microstructure and is controllably releasabl e from the composition, and such that the microfiber has a degree of crystallinity at least 10% greater than that of control polymer prior to wet spinning. For example, the medical condition is at least one selected from the group of: a burn, an abrasion, a laceration, a pathology, a cancer, and an infection. In related embodiments of the method of treating the subject, a first encapsulated therapeutic agent is dexamethasone. In various other embodiments, the method further includes an encapsulated second therapeutic agent. For example, the second therapeutic agent is at least one selected from the group of: a sugar; a carbohydrate; a nucleotide sequence; a protein selected from the group of: a growth factor, an immunoglobulin, an enzyme, and an antibiotic; and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anticoagulant; an anti-tumor; an anti-viral; an anti-bacterial; an anti-mycobacterial; an anti-fungal; an anti-proliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a non- steroidal anti-inflammatory agent (NSAID). For example the nucleotide sequence comprises a vector. For example the vector is a viral or a bacterial vector.

According to various embodiments of the method for treating the subject, the composition has a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold. In related embodiments, the polymers are at least one of poly- 1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP).

Another aspect of the invention provides a kit for treating a subject having a medical condition including a wet spun microfiber composition having at least one polymer such that the composition includes a porous multi-layer polymeric microstructure, and further includes at least one encapsulated therapeutic agent, such that the therapeutic agent is located in an inner core of the microstructure and is controllably releasable from the composition, and such that the microfiber has a degree of crystallinity at least 10% greater than that of control polymer prior to wet spinning; instructions for use; and, a container.

In related embodiments of the kit a first encapsulated therapeutic agent is

dexamethasone. In other embodiments the kit contains a second encapsulated therapeutic agent. For example, the second therapeutic agent is at least one selected from the group of: a sugar; a carbohydrate; a nucleotide sequence; a protein selected from the group of: a growth factor, an immunoglobulin, an enzyme, and an antibiotic; and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anti -coagulant; an anti-tumor; an anti-viral; an antibacterial; an anti-mycobacterial; an anti-fungal; an anti-proliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a non-steroidal anti-inflammatory agent (NSAID). For example, the nucleotide sequences comprises a vector.

According to other embodiments of the kit, the polymers are at least one of poly- 1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP). In related embodiments of the kit the composition has a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

Brief descriptions of the drawings

FIG. 1 panels A-D are scanning electron microscopy photomicrographs of surface and cross-sectional morphology of control and drug loaded PLLA microfibers fabricated by wet spinning. Scale bars: panels A and C, 30 μηι (xlOOO magnification); panels B and D, 30 μπι (xl300 magnification); and (insets) 5 μηι (x5000 magnification).

FIG. 1 panel A columns I and II respectively are: scanning electron microscopy photomicrographs of surface morphology of control and drug loaded PLLA microfibers after fabrication.

FIG. 1 panel B columns I and II respectively are: scanning electron microscopy photomicrographs of cross-sectional morphology of control and drug loaded PLLA microfibers after fabrication.

FIG. 1 panel C columns I and II respectively are: scanning electron microscopy photomicrographs of surface morphology of control and drug loaded PLLA microfibers after eight weeks of incubation in phosphate buffered saline (PBS).

FIG. 1 panel D columns I and II respectively are: scanning electron microscopy photomicrographs of cross-sectional morphology of control and drug loaded PLLA microfibers after eight weeks of incubation in PB S .

FIG. 2 panels A and B are line graphs of drug release as a function of time (eight weeks) of wet spun PLLA microfibers loaded with Dexamethasone (DXM).

FIG. 2 panel A is a graph of percent cumulative release (ordinate) of wet spun PLLA microfibers loaded with 1.0% , 2.4% , and 4.8% (w/w) DXM. Mean ± SEM (Standard error of the mean) are presented.

FIG. 2 panel B is a graph of change in supernatant pH values as a function of degradation time of PLLA microfibers loaded with increasing amounts of DXM in comparison to control microfibers without drug.

FIG. 3 panel A is a scan of representative differential scanning calorimetry (DSC) thermograms of wet spun in vitro 4.8 % (w/w) DXM loaded PLLA microfibers at two week intervals. FIG. 3 panel B compares Hyper DSC thermograms of control and 4.8% (w/w) DXM- loaded microfibers, and free DXM. No melting endotherm was present at 300°C for drug-loaded microfibers. FIG. 4 panel A is a drawing of structure of a semi-crysta!line wet spun polymer. The dark and ordered regions are crystalline and light tangled regions are amorphous.

FIG. 4 panel B is a drawing of formation of new crystalline areas during solvent-induced crystallization (SINC).

FIG. 4 panel C illustrates chemical structural secondary interactions between DXM and PLLA creating a reinforced composite material.

FIG. 5 is a graph of X-ray diffraction pattern traces of control (0%), DXM-loaded PLLA microfibers, and free DXM. No crystalline DXM was detected in any of the microfiber formulations. As the amount of DXM loaded was increased a new crystalline peak appeared in the diffraction pattern.

FIG. 6 panel A is a bar graph of elastic moduli of DXM-loaded PLLA microfibers loaded with 1.0%, 2.4% and 4.8% (w/w) DXM, and control. Values are presented as mean ± SEM. *p <0.05 by one-way analysis of variance (ANOVA)

FIG. 6 panel B is a bar graph of ultimate tensile strength (UTS) of DXM-loaded PLLA microfibers loaded with 1.0%, 2.4% and 4.8% (w/w) DXM, and control. Values are presented as mean ± SEM. *p <0.05 by ANOVA

FIG. 6 panel C is a bar graph of strain at failure of DXM-loaded PLLA microfibers loaded with 1.0%, 2.4% and 4.8% (w/w) DXM, and control. Values are presented as mean ± SEM. *p <0.05 by ANOVA

FIG. 7 panel A is a graph of changes in UTS of DXM-loaded PLLA microfibers loaded with 1.0%, 2.4% and 4.8% (w/w) DXM as a function of time. Values are presented as mean ± SEM. *p <0.05 by ANOVA

FIG. 7 panel B is a graph of comparison of UTS of DXM-loaded PLLA microfibers and control microfibers at eight weeks of incubation in PBS. Values are presented as mean ± SEM. *p <0.05 by ANOVA

FIG. 7 panel C is graph of changes in strain at failure of DXM-loaded PLLA microfibers loaded with 1.0%, 2.4% and 4.8% (w/w) DXM as a function of time. Values are presented as mean ± SEM. *p <0.05 by ANOVA FIG. 7 panel D is graph of comparison of elastic moduli of DXM-loaded PLLA microfibers and control microfibers at eight weeks of incubation in PBS. Values are presented as mean ± SEM. *p <0.05 by ANOVA FIG. 8 panels A and B are polarized light photomicrographs of 4.8% (w/w) DXM-loaded

PLLA microfibers. Scale bars (panels A and B) 50 μιη; (insets) 10 um.

FIG. 8 panel A is an axial cross-section.

FIG. 8 panel B is an orthogonal cross-section. FIG. 9 is a trace of an X-ray diffraction pattern of control PLLA microfibers fabricated using a residence time of 1.5 hours.

FIG. 10 panels A-L are scanning electron micrographs of PLGA microfibers as a function of protein loading. Scale bar for panels A, D, G and J is as in panel A, 50 μπι (35 Ox magnification). Scale bar for panels B, E, H and K is as in panel B, 30 μιη (1 OOOx

magnification). Scale bar for panels C, F, I and L is as in panel C, 50 μι^η (600x magnification). Insert scale bar for panels B, C, E, F, H, 1, and L is 3 μηι (5000x magnification).

FIG. 10 panel A: a cross-sectional morphology of control PLGA microfibers.

FIG. 10 panel B: a cross-sectional morphology of an individual control PLGA microfiber.

FIG. 10 panel C: a surface morphology of an individual PLGA microfiber.

FIG. 10 panel D: a cross-sectional morphology of PLGA microfibers loaded with insulin

(INS).

FIG. 10 panel E: a cross-sectional morphology of an individual PLGA microfiber loaded with INS.

FIG. 10 panel F: a surface morphology of an individual PLGA microfiber loaded with

INS.

FIG. 10 panel G: a cross-sectional morphology of PLGA microfibers loaded with lysozyme (LZ).

FIG. 10 panel H: a cross-sectional morphology of an individual PLGA microfiber loaded with LZ.

FIG. 10 panel I: a surface morphology of an individual PLGA microfiber loaded with

LZ.

FIG. 10 panel J: a cross-sectional morphology of PLGA microfibers loaded with bovine serum albumin (BSA). FIG. 10 panel K: a cross-sectional morphology of an individual PLGA microfiber loaded with BSA.

FIG. 10 panel L: a surface morphology of an individual PLGA microfiber loaded with

BSA.

FIG. 11 panels A-L are scanning electron micrographs of PLLA microfibers loaded with proteins. Scale bar for panels A, D, G and J is as in panel A, 50 μηι (350x magnification). Scale bar for panels B, E, H and K is as in panel B, 30 μπι (lOOOx magnification). Scale bar for panels C, F, I and L is as in panel C, 50 μιη (600x magnification). Inset scale bar for panels B, C, E, F, H, I, K and L is 3 μηι (5000x magnification).

FIG. 1 1 panel A: a cross-sectional morphology of control PLLA microfibers.

FIG. 11 panel B: a cross-sectional morphology of an individual control PLLA microfiber.

FIG. 1 1 panel C: a surface morphology of an individual control PLLA microfiber.

FIG. 1 1 panel D: a cross-sectional morphology of PLLA microfibers loaded with INS.

FIG. 1 1 panel E: a cross-sectional morphology of an individual PLLA microfiber loaded with INS.

FIG. 1 1 panel F: a surface morphology of an individual PLLA microfiber loaded with

INS.

FIG. 1 1 panel G: a cross-sectional morphology of PLLA microfibers loaded with LZ.

FIG. 11 panel H: a cross-sectional morphology of an individual PLLA microfiber loaded with LZ.

FIG. 11 panel I: a surface morphology of an individual PLLA microfiber loaded with

LZ.

FIG. 1 1 panel J: a cross-sectional morphology of PLLA microfibers loaded with BSA.

FIG. 1 1 panel K: a cross-sectional morphology of an individual PLLA microfiber loaded with BSA.

FIG. 1 1 panel L: a surface morphology of an individual PLLA microfiber loaded with

BSA.

FIG. 12 is a graph of thermal analysis of wet spun microfibers. The curves are representative DSC thermograms of: (i) control PLGA microfibers after fabrication, (ii) control PLGA microfibers after fabrication and incubation in phosphate buffered saline (PBS) for 63 days, (iii) control PLLA microfibers after fabrication, and (iv) control PLLA microfibers after fabrication and incubation in PBS for 63 days. FIG. 13 panels A and B are: X-ray diffractograms of control (no protein), INS, LZ, and BS A- loaded PLLA microfibers.

FIG. 13 panel A: diffractograms obtained after fabrication, and FIG. 13 panel B:

diffractograms obtained at 63 days incubation.

FIG. 14 panels A-D are bar graphs of mechanical analyses of wet spun PLGA and PLLA microfibers as a function of protein-loading and polymer type. Mean ± S.D (standard deviation), are presented (n = 5). Symbols *, #,†: significant in one-way ANOVA with Tukey post hoc multiple comparisons tests. Control (blank) and protein loaded PLGA and PLLA microfibers are represented by black (control/blank), dark (INS-loaded), light (LZ loaded), and white (BSA- loaded).

FIG. 14 panel A: load at failure measurements of PLGA and PLLA microfibers.

FIG. 14 panel B: ultimate tensile strength measurements of PLGA and PLLA

microfibers.

FIG. 14 panel C: strain at failure measurements of PLGA and PLLA microfibers.

FIG. 14 panel D: elastic modulus measurements of PLGA and PLLA microfibers.

FIG. 15 panels A-F are line graphs of protein release profiles of cumulative percent release from loaded wet spun PLGA and PLLA microfibers. In panels A and B cumulative percent release is presented as mean ± S.D. for duplicate batches of each formulation (n = 6). In panels C-F, the correlation coefficient (R²) is shown with the exponential regression, y = Ar^x for each time point.

FIG. 15 panel A: cumulative percent release of proteins by PLGA microfibers loaded with INS (squares), LZ (circles) and BSA (triangles) as a function of time.

FIG. 15 panel B: cumulative percent release of proteins by PLLA microfibers loaded with INS (squares), LZ (circles) and BSA (triangles) as a function of time.

FIG. 15 panel C: cumulative percent release of proteins from PLGA fibers after fabrication at day one as a function of protein molecular weight.

FIG. 15 panel D: cumulative percent release of proteins from PLGA fibers after fabrication at day 38 as a function of protein molecular weight. Release kinetics from PGLA fibers was observed to exponentially decrease with increased protein molecular weight.

FIG. 15 panel E: cumulative percent release of proteins from PLLA fibers after fabrication at day one as a function of protein molecular weight. FIG. 15 panel F: cumulative percent release of proteins from PLLA fibers after fabrication at day 38 as a function of protein molecular weight. PLLA fibers encapsulating proteins exhibited sustained release rates independent of molecular weight over the course of 38 days.

FIG. 16 panels A and B are FT-IR spectra of PLGA and PLLA formulations at fabrication (0 day) and at 63 day incubation. Each trace represents the average spectra of 16 scans per sample. FT-IR spectra were offset for clarity.

FIG. 16 panel A: FT-IR spectra of PLGA microfiber formulations of proteins INS, LZ and BSA or blank control (no protein) at fabrication and at 63 days.

FIG. 16 panel B: FT-IR spectra of PLLA microfiber formulations having no protein (blank) or proteins INS, LZ and BSA, or blank control (no protein) at fabrication and at 63 days. FIG. 17 panels A and B are cross-polarized optical micrographs of a PLLA:PLGA (1 : 1) phase separated film (magnification 20x) at 0° and 45° angles respectively.

FIG. 18 panels A and B are graphs of DSC thermograms of PLLA, PLGA and

PLLA:PLGA (1 : 1) microfibers. Two glass transitions were observed, each representative of PLLA and PLGA controls.

FIG. 18 panel A is a graph of DSC thermogram of PLLA:PLGA (1 : 1) microfibers obtained from first heating scan.

FIG. 18 panel B is a graph of DSC thermogram of PLLA:PLGA (1 : 1) microfibers obtained from second heating scan.

FIG. 19 panels A-F are scanning electron micrographs of cross-sectional morphology of binary phase PLLA:PLGA and control PLLA and PLGA microfibers.

FIG. 19 panels A and B are scanning electron micrographs of cross-sectional morphology of PLLA microfibers at low and high magnifications, respectively.

FIG. 19 panels C and D are scanning electron micrographs of cross-sectional morphology of PLGA microfibers at low and high magnifications, respectively.

FIG. 19 panels E and F are scanning electron micrographs of cross-sectional morphology of 1 : 1 PLLA:PLGA microfibers at low and high magnifications, respectively. FIG. 20 panels A and B are fluorescent images of top and side views, respectively of : 1 PLLA.PLGA binary phase composite microfibers encapsulating 0.3% w/w FITC (Fluorescein isothiocyanate)-dextran. FIG. 21 panel A is a line graph of cumulative release kinetics of PLGA, PL LA and

PLLA:PLGA (1 : 1) microfibers loaded with BSA. Fibers with PLLA:PLGA (1 :1) exhibit reduced burst effect in comparison to PLGA and PLLA only fibers.

FIG. 21 panels B is a schematic diagram of phase separated spin dope solutions in the fabrication of binary phase composite microfibers.

FIG. 22 panel A is a schematic representation of a spinning apparatus for wet spinning of polymer into fibers.

FIG. 22 panel B is a photograph of an apparatus for multifilament yarn production (left), a scanning electron micrograph of a multifilament yarn twisted along the longitudinal axis in 'Z' direction (right top), and a schematic representation of a multifilament yarn twisted along the longitudinal axis in 'Z' direction.

FIG. 23 panels A-J are a photograph and scanning electron micrographs of wet spun monofilaments with diverse surface structures.

FIG. 23 panel A is a photographic image of as-spun 3D fiber bundle of 20% (w/v) PLGA extruded into petroleum ether.

FIG. 23 panel B is a scanning electron micrograph of fibers in panel A.

FIG. 23 panel C is a scanning electron micrograph of a post-drawn 7.5% (w/w) PLLAo .94 fiber extruded into a 75:25 ratio of 2-propanol to petroleum ether.

FIG. 23 panel D is a scanning electron micrograph of the fiber in panel C at higher magnification.

FIG. 23 panel E is a scanning electron micrograph of 15% (w/v) PLLA fibers extruded into a petroleum ether coagulation bath.

FIG. 23 panel F is a scanning electron micrograph of 10% (w/v) PLLA fibers extruded into a petroleum ether coagulation bath.

FIG. 23 panel G is a scanning electron micrograph of 7.5% (w/v) PLLA fibers extruded into a 50:50 ratio of petroleum ether to 2-propanol coagulation bath.

FIG. 23 panel H is a scanning electron micrograph of wet spun 7.5% (w/v) PLLA₀.94 fibers extruded into a 50:50 ratio of petroleum ether to 2-propanol coagulation bath. FIG. 23 panel I is a scanning electron micrograph of PLLA/PLGA composite fiber extruded into a petroleum ether coagulation bath.

FIG. 23 panel J is a scanning electron micrograph of PLLA/PLGA composite fiber extruded into a 50:50 ratio of petroleum ether to 2-propanol coagulation bath.

FIG. 24 panels A-H are scanning electron micrographs of post-drawn composite fibers.

FIG. 24 panels A and C are scanning electron micrographs of surface structures of blank and DXM-loaded post-drawn composite fibers, respectively, spun from 10% (w/v) spin dope solution. These fibers were observed to have longitudinal striations with many spherulites.

FIG. 24 panels E and G are scanning electron micrographs of surface structures of blank and DXM-loaded post-drawn composite fibers, respectively, spun from 20% (w/v) spin dope solution. These fibers were observed to have nanoporous surfaces, and fewer spherulites compared to composite fibers in panels A and C.

FIG. 24 panels B and D are scanning electron micrographs of cross-sectional morphology of blank and DXM-loaded post-drawn composite fibers, respectively, spun from 10% (w/v) spin dope solution.

FIG. 24 panels F and H are scanning electron micrographs of cross-sectional morphology of blank and DXM-loaded post-drawn composite fibers, respectively, spun from 20% (w/v) spin dope solution. These fibers were observed to have greater porosity compared to fibers shown in B and D.

FIG. 25 panel A is a line graph of load at failure (strength) of blank wet spun PLLA fibers as a function of polymer concentration.

FIG. 25 panel B is a line graph of strain at failure (ductility) of blank wet spun PLLA fibers as a function of polymer concentration.

FIG. 26 panels A and B are line graphs of modulation of drug release kinetics by composite fibers made with different polymer compositions. The various composite fibers are: PLLA/PLGA (SF-lOa), PLLA/PLGA/PVP (polyvinylpyrolidone; SF-lOb), PLLA/PLLA_{0 94} (SF- 1 Oc), PLLA/PLLA₀.94 PVP (SF-lOd), and PLLA/PVP (SF-lOe).

FIG. 26 panel A is a line graph of drug release profiles of 10 mg wet spun composite monofilaments prepared from 10% (w/v) polymer solution loaded with 2.6% (w/w) DXM.

FIG. 26 panel B is a line graph of drug release profiles of 10 mg wet spun composite monofilaments prepared from 20% (w/v) polymer solution loaded with 1.5% (w/w) DXM. FIG. 27 panels A-C are line graphs of release kinetics of DXM from multifilament yarns.

FIG. 27 panel A is a line graph of drug release profile traces of 10 mg wet spun composite monofilaments prepared from 10% (black circles) and 20% (open circles) (w/v) polymer solutions loaded with 2.8 and 2.4% (w/w) DXM, respectively. Mean ± S.D. are represented.

FIG. 27 panel B is a line graph of multiple theoretical predictions (traces) of drug release kinetics of DXM-loaded multifilament yarns as a function of composition of single monofilament. Each monofilament is a 6-ply yarn combination made of formulations Ά', which is a 10% (w/v) polymer composite, and 'B', which is a 20% (w/v) polymer composite. From top to bottom the monofilament compositions are: 1A + 5B, 5A + IB, 4A + 2B, 3A_bi_ank + 3B, and 3A + 3B_bumk-

FIG. 27 panel C is a plot of drug release profiles of 10 mg 6-ply multifilament yarns produced by 'Z' twisting 4 monofilaments of formulation 'A' and 2 filaments of formulation 'B' (black squares) in comparison to the predicted release calculated using equation 7.

FIG. 28 is a bar graph of biological activity of eluted dexamethasone from multifilament yarns. Dexamethasone treatment inhibits proliferation of human aortic valve interstitial cells (hVICs). Cell number was determined after treatment with 10^"7 mol L^"1 DXM for 72 hours. DXM eluted from 6-ply multifilament yarns after 1 day and 56 days were observed to have the same biological activity as fresh, unencapsulated drug. Mean ± S.D. are presented; *p < 0.05 compared to control by ANOVA.

FIG. 29 panels A, B, and C are scanning electron micrographs, respectively of braided, woven, and complex geometry knitted wet spun multifilament yarns.

FIG. 30 is a drawing of four different levels of hierarchy in the design of therapeutic biomedical textiles. A bottom-up approach (left to right) is demonstrated for fabrication of wet spun filaments as simple building blocks (micro) for the formation of macro-level scaffolds.

Detailed description

Drug loaded micron-scale fibers and methods for their production provided herein are useful for a variety of applications, for example, surgical reconstruction. Local, tunable drug diffusion is useful for constructing anti-neoplastic or immune-privileged boundaries, for example, in the context of reconstructive surgeries and control of local wound healing, fibrosis, scaring, and injury responses including calcification. In vascular reconstructions, modulation of anastamotic healing offers potential to improve patency rates for microvascular repairs, and their presence in blood vessels opens the possibility for prolonged, regional, systemic delivery of therapeutic molecules via the circulation for targeted downstream effects.

Polymeric fibers have medical applications as surgical sutures, dialysis devices, therapeutic implants, wound dressings and tissue engineering (TE) scaffolds. Advances in polymer and drug delivery sciences have led to the evolution of engineered fibers for use as drug delivery vehicles. Design of pharmacologically active fibers has increased (Shibuya et al.

Laryngoscope 2003 ; 1 13( 1 1): 1870-1884; Kim et al. J Control Release 2004;98(l):47-56; Zurita et al. Macromol Biosci 2006;6(9):767-775; Yilgor et al. Biomaterials 2009;30(21):3551-35). Drug eluting fibers have the potential to be knitted, woven, or braided into biotextiles for the release of a multitude of therapeutics with micron-scale accuracy (Tuzlakoglu et al. Tissue Eng Part B Rev 2009; 15(l): 17-27).

Polymer fiber delivery systems by impregnating therapeutics into the core of hollow fibers, entrapping therapeutics within fibers, and chemically crosslinking or adsorbing therapeutics to the surfaces of fibers have been attempted. High surface area to volume ratio of fibers is advantageous for mass transfer and efficient drug release. Release of a model protein, bovine serum albumin (BSA), from the hydrogel cores of co-extruded wet spun PLLA fibers fabricated by Crow et al. exhibited sustained delivery up to eighty days in vitro (Crow et al. Biopolymers 2006; 81(6): 419-427). Sustained delivery of a model hydrophilic anti-cancer drug, 5-fluorouracil, was obtained by impregnating the drug within wet spun PLLA fibers for up to twenty-one days in vitro (Gao et al. J Control Release 2007; 1 18(3): 325-332). Jung et al.

showed that the delivery of cell-permeable gene complexes from PLLA scaffolds improved the transfection of stem cells attached to the surfaces of fibers in comparison to bolus delivery strategies (Jung et al. J Control Release 201 1 ; 152(2): 294-302).

Little is known about the effects of drug incorporation on mechanical integrity of the fibers (Chang et al. J Biomed Mater Res A 2008; 84(1): 230-237; Mack et al. J Control Release 2009; 139(3): 205-211 ; Rissanen et al. J Appl Polym Sci 2010;1 16(4):2174-2180; Williamson et al. Tissue Eng 2006; 12(1):45-51). "Smart" fiber delivery systems are needed that are multifunctional, and provide both physical and pharmaceutical support.

An aspect of the invention provides a composition for delivering a therapeutic agent including: a multi-layer polymeric micro structure including the therapeutic agent, such that the therapeutic agent is located or compartmentalized in an inner core of the microstructure and is characterized by controllable release from the composition. In an embodiment of the composition, the microstructure comprises poly-1 -lactic acid (PLLA) and poly-lactic-co- glycolide (PLGA). In an embodiment of the composition, the composition is porous. In various embodiments of the composition, the therapeutic agent includes at least one selected from the group of: a drug, a protein, a sugar, a carbohydrate, and a nucleotide sequence. In a related embodiment of the composition, the protein includes at least one of the group selected from: a growth factor, an immunoglobulin (antibody), an enzyme, and an antibiotic. In an embodiment of the composition, the nucleotide sequences include a vector. In an embodiment of the composition, the vector comprises a viral vector or a bacterial vector. In an embodiment of the invention, the therapeutic agent includes dexamethasone. In an embodiment of the composition, the therapeutic agent includes a glycoprotein such as a Nog (Noggin) protein.

In an embodiment of the composition, the composition includes at least one selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

An aspect of the invention provides a method of producing a binary phase composition including: mixing a plurality of polymers with a solvent to form a resulting polymer\solvent material; and wet spinning the material by phase inversion, thus producing the binary phase composition. In an embodiment of the method, prior to mixing the plurality of polymers with the solvent, the method includes contacting the plurality with a therapeutic agent.

In an embodiment of the method, the therapeutic agent comprises at least one of the group selected from: a drug, a protein, a sugar, a carbohydrate, and a nucleotide sequence. In an embodiment of the method, the therapeutic agent is at least one of the group selected from: anticoagulant, anti-tumor, anti-viral, anti-bacterial, anti-mycobacterial, anti-fungal, anti- proliferative, anti-inflammatory, anti-apoptotic, immunosuppressant, and pro-apoptotic. In a related embodiment of the method, the anti-inflammatory is selected from: a steroid and a nonsteroidal anti-inflammatory agent (NSAID). For example the steroid is selected from the group of: a cortisone compound for example a dexamethasone: and a sex-related hormone.

In an embodiment of the method, the solvent includes at least one selected from the group of: chloroform, dichloromethane, ethyl acetate, diethyl ether, acetic acid, hexane, ethanol, methanol, acetone, tetrahydrofuran, toluene, dimethyl sulfoxide, acetonitrile, and a combination thereof. In various embodiments of the method, wet spinning includes loading the material into a syringe, and dispensing the material into a coagulation bath including a non-solvent for example petroleum ether. In an embodiment of the method, the coagulation bath includes petroleum ether. In various embodiments of the method, a difference between the solubility parameter of the solvent and the solubility parameter of the non-solvent affects the rate of solidification and the degree of crystallinity of the microstructure. The difference between the solubility parameter of the solvent and the solubility parameter of the non-solvent is selected from one of the following: less than about 12 units, less than 10 units, less than 9 units, less than 8 units, less than 7 units, less than 6 units, less than 5 units, less than 4 units, less than 3 units, less than 2 units and less than 1 unit. In various embodiments, the difference is less than about 2-4 units, less than about 4-6 units, less than about 6-8 units, or less than about 8-10 units. In various embodiments, the difference is varied to modulate rate of crystallization of the composition. In a related embodiment, the plurality of polymers includes a polymer matrix or a composite material. For example, the polymer matrix is bioabsorbable.

An aspect of the invention provides a method of treating a subject having a medical condition including: contacting the subject with a composition including a multi-layer polymeric microstructure including a therapeutic agent, such that the therapeutic agent is located in an inner core of the microstructure and is characterized by controllable release from the composition.

In a related embodiment of the method, the composition includes at least one selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold. In various embodiments of the method, the medical condition is at least one selected from the group of: a burn, a cut, an abrasion, a laceration, a pathology, a cancer, and an infection, in a related embodiment of the method, the microstructure includes poly- 1 -lactic acid (PLLA) and poly-lactic-co-glycolide (PLGA). In an embodiment of the method, the therapeutic agent comprises at least one of the group selected from: a drug, a protein, a sugar, a carbohydrate, and a nucleotide sequence. In various embodiments of the method, the protein includes at least one of the group selected from: a growth factor, an immunoglobulin, an enzyme, and an antibiotic. In an embodiment of the method, the nucleotide sequence includes a vector. In a related embodiment of the method, the vector includes a viral vector or a bacterial vector. In an embodiment of the method, the therapeutic agent is a corticosteroid for example a dexamethasone.

An embodiment of the invention provides a kit for treating a subject in need of medical treatment including: a composition for delivering a therapeutic agent including a multi-layer polymeric microstructure including the therapeutic agent, such that the therapeutic agent is located in an inner core of the microstructure and is characterized by controllable release from the composition; instructions for use; and, a container.

In related embodiments of the kit, the microstructure includes poly- 1 -lactic acid (PLLA) and poly-lactic-co-glycolide (PLGA). In an embodiment of the kit, the therapeutic agent includes at least one selected from the group of: a drug, a protein, a sugar, a carbohydrate, and a nucleotide sequence. In various embodiments of the kit, the protein includes at least one selected from the group of: a growth factor, an immunoglobulin, an enzyme, and an antibiotic. In an embodiment of the kit, a nucleotide sequence includes a vector. In a related embodiment of the kit, the vector includes a viral vector or a bacterial vector. In an embodiment of the kit, the therapeutic agent includes dexamethasone. In an embodiment of the kit, the composition includes at least one selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

The composition and methods described herein are provided for drug encapsulation and processing conditions that affect the mechanical integrity of microfibers, including producing microfibers that perform a surgical mechanical function and simultaneous drug delivery. Drug- drug and drug-polymer interactions were evaluated herein using wet spun PLLA microfibers loaded with 1.0, 2.4, and 4.8% (w/w) DXM. In the spinning of semi-crystalline polymers, the crystalline regions of the polymer contribute to strength. PLLA was selected herein for wet spinning because of its material properties. PLLA contains ester groups and DXM contains two carbonyl and three hydroxyl groups. Without being limited by any particular theory or mechanism, hydrogen bonding occurs between the carbonyl oxygen atoms in PLLA chains and the hydroxyl hydrogen atoms in DXM. Hydrogen bonding during incubation is enabled as the amorphous regions of the polymer become more mobile.

Wet spinning is a technique that is here applied to drug delivery technologies with advantages of ambient temperatures manufacture. Wet spinning is initiated by dissolving a polymer in solvent. The dissolved solution is extruded through a spinneret and into a non- solvent coagulation bath. The solvent is miscible with the non-solvent, and the polymer in solution is not, and a continuous polymer stream precipitates into a solid filament (Gupta et al. 2007 Prog Polym Sci 32(4):455-482). Wet spinning is used also for encapsulation of water- soluble drugs since both the solvent and coagulant can be non-aqueous, which produces a hydrophobic environment, thereby significantly reducing the leaching of water-soluble drugs from during the encapsulation process. A broad range of bioactive agents including antibiotics, heparin, proteins, growth factors, genes, and even viruses have been successfully wet spun into fibers for many biomedical applications (Blaker et al. Biomaterials 2004 25(7-8): 1319-1329; Pasternak et al. Int J Colorectal Dis 2008; 23(3): 271-276; Hirano et al. J Biomed Mater Res 2001 ;56(4): 556-561 ; Cronin et al J Biomed Mater Res A 2004;69(3): 373-381 ; Hwang et al. Langmuir 2008; 24(13): 6845-6851 ; Crow BB et al. Biopolymers 2006; 81(6): 419-427; Jung et al. J Control Release 2011 ; 152(2): 294-302; Chiang et al. Adv Mater 2007; 19(6): 826-827).

PLGA and PLLA are biodegradable materials and have FDA approval for many medical applications. Three proteins of different molecular weights, insulin (5.8 kDa), lysozyme (14.3 kDa), and bovine serum albumin (66.0 kDa) were encapsulated in PLGA and PLLA microfibers to analyze the effect of protein molecular weight and polymer type on release kinetics and intrinsic material properties of wet spun microfibers. The optimal protein loading of 2% (w/w) was determined based on the maximum amount of BSA that was loaded into 20% (w/v) polymer concentrations without disrupting the continuous formation of fibers. Therapeutics with molecular weights higher than BSA are envisioned to be incorporated into fibers by increasing the concentration of the spin dope, increasing the molecular weight of the polymer, or by decreasing the theoretical loading of the protein. Other polymers are fabricated into wet spun microfibers by selecting the appropriate solvents and nonsolvents.

Fibers prepared from cryogenic emulsions were observed to have non-circular skin-core structure consistent with the wet extrusion process. Fiber shape is in part a result of solvent and nonsolvent counter-diffusion. If the rate of solvent diffusing out is higher than the rate of nonsolvent diffusing in, the fiber structure collapses and non-circular shapes such as lobed 'kidney' and 'dog bone' are formed (Sobhanipour P et al., Thermochim Acta, 201 1 ;518: 101- 106). Rapid surface coagulation during phase inversion leads to the entrapment of solvent and nonsolvent within the precipitating microfilament (Rissanen M et al., J Appl Polym Sci, 2008;1 10:2399-2404). Porous structure is formed by evaporation of solvent and nonsolvent after microfilament solidification. PLGA and PLLA microfibers prepared using methods herein have significantly less voids than poly(L,D-lactic acid) fibers prepared from water-in-oil (W/O) emulsions by Rissanen et al. (Rissanen M et al., J Appl Polym Sci, 2010; 1 16:2174-2180). Pores were attributed to air bubbles from emulsion formation and rapid phase separation during filament precipitation in the spin bath. Cryogenic emulsion process described herein reduces the potential for large voids by removing the water phase from the primary W/O spin dope. Void formation was reduced by encapsulating drugs as solid particles within wet spun microfibers (Gao H et al, J Control Release, 2007; 1 18:325-332), and hydrophilic drugs encapsulated were not micronized and thus were characterized by large particles imbedded within the fibers and on the surfaces of fibers.

Thermal analyses of fibers (t = 0) showed that the glass transition temperatures of PLGA formulations did not substantially decrease with protein loading, nor did the relative crystallinity of PLLA formulations. These results indicated that the long duration (-1.5 h) of the fibers in the nonsolvent bath induced solvent-induced crystallization (SINC). An increase in glass transition temperature was observed among formulations with incubation (t = 63). The increase in glass transition temperature was due to increased amorphous chain mobility at the incubation temperature (37°C) and potential formation of ordered structures. Similar thermal induced crystallization with incubation were observed in the preparation of DXM-loaded PLLA microfibers.

X-ray diffraction studies were used to further evaluate the molecular morphology of wet spun fibers. The PLLA samples in Examples herein, including blank fibers, had similar x-ray diffraction patterns, and SINC was persistent despite the presence of the encapsulated protein. In the degraded PLLA formulations, the peak at 24.4° disappeared at 63 day incubation due to hydration and re-orientation of the polymer. The new sharp peak that appeared at 16.5° corresponded to amorphous chain restructuring of low molecular weight chains in the polymer with incubation. Re-ordering of shorter oligomers and strong secondary interactions between proteins and polymers further contributed to this effect.

The molecular weight of the protein and polymer structure also influenced the physical properties of wet spun microfilaments. Both polymers exhibited significantly reduced tensile strength with increased protein molecular weight and PLLA microfibers loaded with BSA were observed to have the greatest tensile loss. Without being limited by any particular theory or mechanism, the reduced tensile strength is in part a result of the differences in polymer structure between PLGA and PLLA. PLLA is a semi-crystalline polymer whereas PLGA is amorphous. In general, the amorphous regions of a polymer require less force to deform in comparison to the crystalline regions. At high protein loading with BSA, the material properties of PLGA and PLLA microfibers were significantly weakened possibly due to protein particles acting as material defects within the polymer lattice during elastic deformation. The fusion of BSA particles also contributed to these results, as was observed with scanning the electron micrographs (FIGs 10 and 11). However, BSA-PLGA microfibers were observed to have higher tensile strengths and elongations until failure than BSA-PLLA formulations. The amorphous PLGA appeared to reduce the tensile loss and embrittlement of BSA-loaded microfibers to a lesser degree relative to BSA-loaded PLLA formulations. Differences between amorphous PLGA and semi-crystalline PLLA were reflected in elastic moduli. PLGA microfibers had similar resistance to deformation regardless of protein loading. PLLA is semi-crystalline, therefore low protein loading provided slightly increased resistance to deformation likely due to protein particles reinforcing the amorphous regions of the polymer. Protein-polymer interactions at the molecular level also affect the physical properties of the microfiber. Strong secondary interactions between proteins and polymers, such as hydrogen bond formation and ionic interactions helped to maintain the material properties of polymeric microfibers with INS and LZ loading. Encapsulation of small molecules decreased wet spun fiber strength and ductility (Mack BC et al., J Control Release, 2009; 139:205-21 1 ; Williamson MR et al., Biomaterials, 2004;25:5053-5060; Chang HI et al., J Biomed Mater Res Part A, 2008;84:230-237). In examples herein, the effect of three proteins on protein-loaded microfibers was evaluated, and protein particle size was observed to play a critical role in the tensile strength of polymeric wet spun microfibers.

Protein release from wet spun microfibers was found to depend on protein molecular weight. These data show that the amount of protein released is controlled primarily by diffusion. Three phases were seen in the release of INS from PLGA and PLLA formulations, indicating that some degradation was occurring. In the degradation of polyesters, random cleavage of ester linkages along the polymer backbone breaks long polymer chains into short fragments that may not be water-soluble. A reduction in the molecular weight increases hydrophilicity, and additional release of INS with little polymer degradation was observed by DSC and FT-IR analyses, due to slow degradation of wet spun aliphatic polyesters with similar molecular weights (Crow BB et al., Tissue Engineering, 2005 ; 1 1 : 1077-1084; Nelson KD et al., Tissue Engineering, 2003;9: 1323-1330).

The dynamic mechanical properties of drug-eluting wet spun fibers in vitro were evaluated in Examples herein. Dexamethasone (DXM), a synthetic anti-inflammatory glucocorticoid was used in compositions, methods and kits herein as the model hydrophobic drug. Local delivery of DXM from microspheres has been shown to reduce cellular immune response to medical implants (Hickey et al. J Biomed Mater Res 2002; 61(2): 180-187; Patil SD et al. Diabetes Technol Ther 2004; 6(6): 887-897; Barcia et al. Exp Eye Res 2009; 89(2):238- 245). Therefore, DXM elution as described herein is beneficial for reducing unwanted inflammatory responses of fibrous implants. Examples herein analyzed the drug-polymer interactions and the effects of DXM loading and release on the material properties of wet spun PLLA fibers.

Wet spinning by phase inversion was found in Examples herein to produce durable dexamethasone-eluting PLLA fibers that had prolonged release of drug and retained high mechanical strength. It was observed that these microstructures are useful as a fiber delivery system. Sustained release of drug from wet spun PLLA fibers has been a unfulfilled goal of researchers and industry, and Examples herein are the first to report a stable and controlled fiber based delivery system with linear release of <28% total encapsulated drug after eight weeks in vitro. Methods and systems herein produced fibers that have been woven or knitted into tissue- engineering TE scaffolds. Without being limited by any particular theory or mechanism of action, it is here envisioned that dexamethasone included in microstructure compositions (e.g., fibrous scaffolds) reduced innate immune response and resulted in the mechanical support suitable for tissue integration. Furthermore, the addition of an inert hydrophobic molecule, such as dexamethasone, into wet spun fibers resulted in the mechanical properties of microfilaments, and decreased the burst release of hydrophilic therapeutics. A stable delivery system using a porous polymeric microstructure composition was obtained that is physically

manipulatable/easily shaped, and delivers controlled release of therapeutics and maintains mechanical strength. The compositions, methods, and kits using a multi-layer polymeric microstructure are useful for many therapeutic applications including regenerative medicine and tissue engineering. Fabrication of biologically active fibers for integrating into existing biomedical implants, for instance, local, controlled delivery of anti-inflammatory drugs to surrounding tissues would function to reduce unwanted inflammatory cell infiltrates and increase the longevity of implanted biomaterials. A bench-top technique for the scale up of monofilaments into multifilament yarns was developed to enhance the handling capabilities of wet spun filaments and to demonstrate the ability to tune drug release kinetics.

The wet spinning technique in Examples herein is a versatile method for the production of continuous micron-sized fibers. The surface topography of wet spun filaments was manipulated by altering wet spinning parameters. In general, quick quenching yields fibers with smooth surfaces (Xiang HB et al., Macromol Res, 201 1 :19:645-653). Therefore, it was possible to create fibers with micrometer-range features such as grooves, ridges, and spherical protrusions by decreasing the polymer solution concentration, slowing the counter-diffusion of solvent and nonsolvent, altering the residence time in the coagulation bath, or simply applying tension through solution- and post-drawing methods. Altering the surface topography of wet spun fibers is beneficial for host tissue integration and wound healing by enhancing contact guidance and cellular attachment (Cao H et al., J Biomed Mater Res Part A, 2010;93: 1151- 1 159).

To demonstrate the ability to modulate drug release from wet spun filaments, fibers with varying polymer compositions were produced, including compositions having excipients such as PVP, and by using different polymer concentrations. Drug-loaded formulations were as-spun, solution-drawn or post-drawn, depending on the rate of precipitation and counter-diffusion of solvent and nonsolvent. Each of these processes was developed to determine whether wet spun fibers can be used to design hybrid devices, ones that perform a mechanical function and simultaneously deliver drugs. Examples herein show the complexity of designing a wet spun delivery system and the scale up of monofilaments into multifilament yarns.

Overall, DXM release of wet spun filaments of similar polymer concentration was dependent on the hydrophilicity of the composite formulation. For 10% (w/v) solution-drawn composite formulations, the highest drug release rates were achieved by adding PLGA and PVP. Since PVP is a water-soluble polymer, increased drug release was attributed to pore formation from the solubilization of PVP in buffer solution. After the initial burst period, the release of drug from solution-drawn 10% (w/v) formulations was favored by the swelling capability of the fibers. As-spun formulations with 20% (w/v) solutions containing PLLA and PLGA also showed the release of DXM was dependent on the hydrophilicity of the filaments. A substantial difference in drug release kinetics of 20% (w/v) composite fibers was achieved with the addition of 50% PLGA content. The ability to prolong the release of drug was also apparent by increasing the polymer concentration from 10% to 20% (w/v).

Post-drawn composite formulations selected for multifilament yarn production were observed to have increased DXM release with increased PLGA and PVP addition, despite differences in overall polymer concentration. Fibers spun from 10% (w/v) solutions were observed to have reduced DXM burst release in comparison to fibers spun from 20% (w/v) solutions. For these formulations, variations in DXM release kinetics were attributed to processing conditions. The PVP content in 10% (w/v) formulations was less than that in 20% (w/v) formulations, which resulted in the decrease in the initial burst release of DXM. Nano- sized pores on the surfaces of 20% (w/v) composite solutions also modulated drug release by increasing buffer penetration. Surface pore formation is a result of the rapid evaporation of solvent entrapped within the precipitating fiber during winding. Filaments spun from 10% (w/v) solutions also displayed a dense cross-sectional morphology. Cross-sectional pores are caused by the entrapment of solvent during spin dope solution precipitation. Composite fibers from 10% (w/v) solutions were observed to have less polymer concentration to block the counter- diffusion of solvent and nonsolvent, resulting in significantly less cross-sectional porosity in comparison to 20% (w/v) solutions. While both 10% and 20% (w/v) composite fibers experienced similar forces during drawing, 10% (w/v) solutions were observed to have less resistance to deformation as judged by the grooved surface topography.

The therapeutic range of DXM delivery to treat post-operative inflammation from various biomaterials has been evaluated in a number of animal models (Barcia E et al., Exp Eye Res 2009;89:238-245; Dang TT et al., Biomaterials, 201 1 ;32:4464-4470;39; Hickey T et al, J Biomed Mater Res, 2002;61 : 180-187; Patil SD et al. Diabetes Technol Ther, 2004;6:887-897; Selvam S et al, Biomaterials, 201 1 ;32:7785-7792). Hickey et al. evaluated the potential of localized DXM delivery from PLGA microspheres to suppress inflammatory responses to cotton threads implanted subcutaneously in rats (Hickey T et al, J Biomed Mater Res, 2002;61 : 180- 187). They showed DXM treatment was most effective when delivered as a burst release followed by a slow release (3-30 μg day-1) over the course of 30 days in vivo. Microspheres and hydrogel composites have been evaluated by others for the continuous release of DXM to treat biomaterial-driven inflammation and release rates of 0.17-7.2 μg day^*1 were found to be effective in modulating host immune responses (Barcia E et al, Exp Eye Res 2009;89:238-245 ; Patil SD et al. Diabetes Technol Ther, 2004;6:887-897). Real-time inflammatory response to biomaterial implants has been investigated using non-invasive fluorescence imaging techniques (Dang TT et al, Biomaterials, 2011 ;32:4464-4470; Selvam S et al, Biomaterials, 201 1 ;32:7785- 7792). These results showed that local delivery of DXM from PLGA microspheres was successful in significantly reducing biomaterial-driven inflammatory responses continuing until one month in vivo.

To engineer fibers with various release profiles of biologically active drug post-drawn filaments were scaled up into 6-ply multifilament yarns. Yarn production did not affect the predicted release of DXM from individual filaments (FIG. 27 panel C). The biological activity of eluted drug also was preserved until 56-day incubation. Drug-eluting yams were formed from a combination of several types of individual monofilaments for the prolonged release of a multitude of therapeutics with retained biological activity. Drug release profiles were further tuned to meet specific clinical needs by altering the polymer composition, molecular weight and concentration as previously discussed. Additionally, wet spun fibers were used as conduits for the long-term delivery of other biologically active drugs and/or proteins.

The tensile properties of post-drawn monofilaments and 'Z' twisted multifilament yarns were characterized to evaluate their use as multifunctional delivery systems. Increasing the polymer concentration did not result in increased tensile properties as was found with blank fibers spun from PLLA only (FIG. 25). Polymer chains align in the direction of shear flow during spinning (Graessley WW, J Chem Phys, 1965, 43:2696-2703). Dilute polymer solutions have less polymer chains and more mobility to align in the direction of flow, promoting a higher degree of chain orientation along the fiber axis. Thus, it is possible that the molecular orientation and crystalline morphology of post-drawn 10% and 20% (w/v) fibers are different. Post-drawing increases the degree of orientation, density and fiber crystallinity (Arbab S et al., Polym Bull, 2011 ;66: 1267-1280; Williamson MR and Coombes AG, Biomaterials, 2004;25:459-465).

Longitudinal striations on the surfaces of 10% (w/v) fibers indicated that post-handling processes also influenced molecular orientation during fiber drawing. When fibers are under tension, such as in filament drawing, the amorphous units of the polymer begin to unfold, allowing for a reduction in free volume and amorphous polymer chain alignment until fracture.

Monofilaments spun from 10% (w/v) solutions were observed to have similar ductility as compared to 20% (w/v) solutions. One would expect increased polymer concentration to lead to increased strain at failure (FIG. 25). However, the cross-sectional morphology of composite fibers was very different, unlike the as-spun fibers made from PLLA only. The residence time of post-drawn composite fibers prepared for multifilament yarn fabrication was ~6 min, whereas as-spun PLLA fibers remained in the coagulation bath until the spin dope was extruded, yielding a residence time of aboutl .5 hrs. Reduced residence time in the coagulation bath leads to more porous fibers by decreasing the growth rate of voids, and available time for growth during phase separation. Drug encapsulation did not considerably affect the material properties of monofilaments. The United States Pharmacopeia (USP) determines the standards specifying test procedures and product specifications for surgical sutures including the knot-pull tensile strength for bioresorbable sutures. The average maximum tensile stress for DXM-loaded multifilament yarns was approximately 450 mN, or 0.05 kgf, equivalent in strength to a 9-0 (0.030-0.039 mm diameter) absorbable synthetic suture (United States Pharmacopeia and National Formulary (USP 34-NF 29), United States Pharmacopeial Convention, Rockville, MD, 2010). The multifilament yarns are much larger in diameter than 9-0 sutures and their handling capabilities demonstrate the potential to hybridize wet spun fibers with existing surgical sutures to meet tensile strength specifications. Using textile industry embroidery technologies, it is possible to weave, braid, or knit drug-eluting yarns with or around existing sutures. Since there was much strain left in the fibers after the wet spinning process, it is envisioned that mechanical properties of drug-eluting filaments can be improved through industrial-scale post-drawing techniques. Multifilament drug-eluting yarns described in Examples herein stretched over 100% of their initial length before failure. Mechanical stretching is envisioned to be useful to decrease the ductility of yarns and increase tensile strength. Stretching is similarly envisioned to be applied to individual fibers prior to multifilament yarn formation.

A multifilament yarn with tunable DXM release kinetics that is controllable through the combination of constituent monofilaments is described. The spatiotemporal release of therapeutics is further controlled by the location of specific yarns within a 3-dimensional biomedical implant. Examples herein show that the encapsulation of dexamethasone within wet spun fibers does not weaken mechanical strength or lead to fiber embrittlement. Drug-eluting yarns were not as strong as conventional melt spun sutures, and were capable of physical manipulation and have the potential to be incorporated into existing biomedical textiles.

Compositions and methods herein demonstrated the feasibility of making micron-scale alterations to the surface topography of wet spun fibers by applying stretch and varying wet spinning processing conditions. The potential to alter the surface topography of monofilaments, tune the release kinetics of a biologically active therapeutics, and enhance the load bearing strength of wet spun fibers through multifilament twisting, makes wet spun yarns extremely valuable for many biomedical and drug delivery applications (FIG. 30). The diverse release kinetics achieved in Examples herein show an ideal release profile of potential therapeutics to decrease non-specific inflammatory responses to implanted biomaterials.

Methods and compositions herein include a therapeutic agent for example a vector or an antibody. Methods use construction of expression vectors containing a sequence encoding a protein operably linked to appropriate transcriptional and translational control elements. These methods include in vitro recombinant DNA techniques, synthetic techniques and in vivo recombination or genetic recombination. Such techniques are described in Sambrook et al., Molecular Cloning: A Laboratory Manual, Cold Spring Harbor Press, Plainview, NY, 1989.

A variety of commercially available expression vector/host systems are useful to contain and express a protein encoding sequence. These include but are not limited to microorganisms such as bacteria transformed with recombinant bacteriophage, plasmid or cosmid DNA expression vectors; yeast transformed with yeast expression vectors; insect cell systems contacted with virus expression vectors (e.g. , baculovirus); plant cell systems transfected with virus expression vectors (e.g., cauliflower mosaic virus, CaMV; tobacco mosaic virus, TMV) or transformed with bacterial expression vectors (e.g., Ti, pBR322, or pET25b plasmid); or animal cell systems. See Ausubel et al, Current Protocols in Molecular Biology, John Wiley & Sons, New York, NY, 1989. Exemplary viral vectors include adenovirus vectors, lentivirus vectors, adeno-associated virus (AAV) vectors, and helper-dependent adenovirus vectors.

General methodologies for antibody production, including criteria to be considered when choosing an animal for the production of antisera, are described in Harlow et al. (Antibodies, Cold Spring Harbor Laboratory, pp. 93-117, 1988). For example, animals of suitable size such as goats, dogs, sheep, mice, or camels are immunized by administration of an amount of immunogen, such as the intact protein or a portion thereof containing an epitope from human protein, effective to produce an immune response. The technique of in vitro immunization of human lymphocytes is used to generate monoclonal antibodies. Techniques for in vitro immunization of human lymphocytes are described in Inai, et al., Histochemistry, 99(5):335 362, May 1993; Mulder, et al, Hum. Immunol., 36(3): 186 192, 1993; Harada, et al., J. Oral Pathol. Med., 22(4): 145 152, 1993; Stauber, et al., J. Immunol. Methods, 161(2): 157 168, 1993; and Venkateswaran, et al, Hybridoma, 11(6) 729 739, 1992. These techniques can be used to produce antigen-reactive monoclonal antibodies, including antigen-specific IgG, and IgM monoclonal antibodies.

Pharmaceutical compositions

An aspect of the present invention provides pharmaceutical compositions that include a multi-layer polymeric microstructure comprising a therapeutic agent. In related embodiments, the pharmaceutical microiiber composition is formulated sufficiently pure for administration to a human subject, e.g., to an abdomen, an eye, or an appendage of a human subject.

In certain embodiments, the therapeutic agent or agents are selected from the group consisting of growth factors, anti-inflammatory agents, vasopressor agents including but not limited to nitric oxide and calcium channel blockers, collagenase inhibitors, topical steroids, matrix metalloproteinase inhibitors, ascorbates, angiotensin II, angiotensin III, calreticulin, tetracyclines, fibronectin, collagen, thrombospondin, transforming growth factors (TGF), keratinocyte growth factor (KGF), fibroblast growth factor (FGF), insulin-like growth factors (IGFs), IGF binding proteins (IGFBPs), epidermal growth factor (EGF), platelet derived growth factor (PDGF), neu differentiation factor (NDF), hepatocyte growth factor (HGF), vascular endothelial growth factor (VEGF), heparin-binding EGF (HBEGF), thrombospondins, von Willebrand Factor-C, heparin and heparin sulfates, and hyaluronic acid.

In other embodiments, the agent is a compound, composition, biological or the like that potentiates, stabilizes or synergizes the effects of a microfiber on a cell or tissue. In some embodiments, the drug includes without limitation anti-tumor, antiviral, antibacterial, anti- mycobacterial, anti-fungal, anti-proliferative or anti-apoptotic agents. Drugs for inclusion in the microfiber are described in Goodman & Gilman's The Pharmacological Basis of Therapeutics, 9th Ed., Hardman, et al., eds., McGraw-Hill, 1996, the contents of which are herein incorporated by reference herein.

As used herein, the term "pharmaceutically acceptable carrier" includes any and all solvents, diluents, or other liquid vehicle, dispersion or suspension aids, surface active agents, isotonic agents, thickening or emulsifying agents, preservatives, solid binders, lubricants and the like, as suited to the particular dosage form desired. Remington's Pharmaceutical Sciences Ed. by Gennaro, Mack Publishing, Easton, PA, 1995 provides various carriers used in formulating pharmaceutical compositions and known techniques for the preparation thereof. Some examples of materials which can serve as pharmaceutically acceptable carriers include, but are not limited to, sugars such as glucose and sucrose; excipients such as cocoa butter and suppository waxes; oils such as peanut oil, cottonseed oil, safflower oil, sesame oil, olive oil, corn oil, and soybean oil; glycols such a propylene glycol; esters such as ethyl oleate and ethyl laurate; agar; buffering agents such as magnesium hydroxide and aluminum hydroxide; alginic acid; pyrogen-free water; isotonic saline; Ringer's solution; ethyl alcohol; and phosphate buffer solutions, as well as other non-toxic compatible lubricants such as sodium lauryl sulfate and magnesium stearate, as well as coloring agents, releasing agents, coating agents, preservatives and antioxidants can also be present in the composition, the choice of agents and non-irritating concentrations to be determined according to the judgment of the formulator.

Therapeutically effective dose

Methods provided herein involves contacting a subject with a pharmaceutical microfiber composition, for example, administering a therapeutically effective amount of a pharmaceutical composition having as an active agent a multi-layer polymeric microstructure comprising a therapeutic agent, to a subject in need thereof, in such amounts and for such time as is necessary to achieve the desired result. The compositions, according to the method of the present invention are administered using an amount and route of administration effective for treating a subject. The exact dosage is chosen by the individual physician in view of the patient to be treated. Dosage and

administration are adjusted to provide sufficient levels of the active agent(s) or to maintain the desired effect. Additional factors which may be taken into account include the severity of the disease state, e.g., intermediate or advanced stage of AMD; age, weight and gender of the patient; diet, time and frequency of administration; route of administration; drug combinations; reaction sensitivities; and tolerance/response to therapy. Long acting pharmaceutical compositions might be applied at convenient intervals every week, or once every two weeks, month, semi-annually, depending on half-life and clearance rate of the microfiber apparatus.

A therapeutically effective dose refers to that amount of active agent that ameliorates the symptoms or prevents progression of pathology or condition. Therapeutic efficacy and toxicity of active agents can be determined by standard pharmaceutical procedures in cell cultures or experimental animals, e.g., ED50 (the dose is therapeutically effective in 50% of the population) and LD50 (the dose is lethal to 50% of the population) and by release from the microfiber composition. The dose ratio of toxic to therapeutic effects is the therapeutic index, and it can be expressed as the ratio, LD50/ED50. Pharmaceutical compositions which exhibit large therapeutic indices are preferred. The data obtained from cell culture assays and animal studies are used in formulating a range of dosage for human use.

Administration of a source of expression of a protein is administration of a dose of a viral vector or a nucleic acid vector, such that the dose contains for example at least about 50, 100, 500, 1000, or at least about 5000 particles per cell to be treated.

Administration of pharmaceutical compositions

As formulated with an appropriate pharmaceutically acceptable carrier in a desired dosage, the pharmaceutical composition provided herein is administered to humans and other mammals topically such as ocularly (as by a microfiber application), nasally, bucally, orally, rectally, parenterally, intracisternally, intravaginally, or intraperitoneally.

The pharmaceutical composition in various embodiments is administered with inert diluents commonly used in the art such as, for example, water or other solvents, solubilizing agents and emulsifiers such as ethyl alcohol, isopropyl alcohol, ethyl carbonate, ethyl acetate, benzyl alcohol, benzyl benzoate, propylene glycol, 1,3-butylene glycol, dimethylformamide, oils (in particular, cottonseed, groundnut, corn, germ, olive, castor, and sesame oils), glycerol, tetrahydrofurfuryl alcohol, polyethylene glycols and fatty acid esters of sorbitan, and mixtures thereof. Besides inert diluents, the delivered microfiber compositions can also include adjuvants such as wetting agents, and emulsifying and suspending agents. Dosage forms for topical or transdermal administration of a microfiber inventive pharmaceutical composition include ointments, pastes, creams, lotions, gels, powders, solutions, sprays, inhalants, or patches. The active agent is admixed under sterile conditions with a pharmaceutically acceptable carrier and any needed preservatives or buffers as may be required. For example, ocular or cutaneous routes of administration are achieved with aqueous drops, a mist, an emulsion, or a cream. Administration may be therapeutic or it may be prophylactic. The invention includes delivery devices, surgical devices, audiological devices or products which contain disclosed microfiber compositions (e.g., as supplied as a portion of gauze bandages or strips), and methods of making or using such devices or products. These devices may be coated with, impregnated with, bonded to or otherwise treated with a composition as described herein.

Transdermal patches have the added advantage of providing controlled delivery of the active ingredients to the body. Such dosage forms can be made by dissolving or dispensing the compound in the proper medium. Absorption enhancers can also be used to increase the flux of the compound across the skin. The rate can be controlled by either providing a rate controlling membrane or by dispersing the compound in a polymer matrix or gel.

Injectable preparations, for example, sterile injectable aqueous or oleaginous suspensions may be formulated according to the known art using suitable dispersing or wetting agents and suspending agents. The sterile injectable preparation include a sterile injectable solution, suspension or emulsion in a nontoxic parenterally acceptable diluent or solvent, for example, as a solution in 1,3-butanediol. Among the acceptable vehicles and solvents are water, Ringer's solution, U.S.P. and isotonic sodium chloride solution. In addition, sterile, fixed oils are conventionally employed as a solvent or suspending medium, for example a bland fixed oil is employed including synthetic mono- or diglycerides. In addition, fatty acids such as oleic acid are used in the preparation of injectables. The injectable formulations are sterilized, for example, by filtration through a bacterial-retaining filter, or by incorporating sterilizing agents in the form of sterile solid compositions which can be dissolved or dispersed in sterile water or other sterile injectable medium prior to use. In order to prolong the effect of an active agent, it is often desirable to slow the absorption of the agent from subcutaneous or intramuscular injection. Delayed absorption of a parenterally administered active agent is accomplished by dissolving or suspending the agent in an oil vehicle. Injectable depot forms are made by forming microencapsule matrices of the agent in biodegradable polymers such as polylactide- polyglycolide. Depending upon the ratio of active agent to polymer and the nature of the particular polymer employed, the rate of active agent release is controlled. Examples of other biodegradable polymers include poly(orthoesters) and poly(anhydrides). Depot injectable formulations are prepared also by entrapping the agent in liposomes or microemulsions which are compatible with body tissues.

Compositions for rectal or vaginal administration are preferably microfiber suppositories which can be prepared by mixing the active agent(s) of the invention with suitable non-irritating excipients or carriers such as cocoa butter, polyethylene glycol or a suppository wax which are solid at ambient temperature but liquid at body temperature and therefore melt in the rectum or vaginal cavity and release the active agent(s).

Solid dosage forms for oral administration include capsules, tablets, pills, powders, and granules. In such solid dosage forms, the active agent is mixed with at least one inert, pharmaceutically acceptable excipient or carrier such as sodium citrate or dicalcium phosphate and/or fillers or extenders such as starches, sucrose, glucose, mannitol, and silicic acid; binders such as, for example, carboxymethylcellulose, alginates, gelatin, polyvinylpyrrolidinone, sucrose, and acacia; humectants such as glycerol; disintegrating agents such as agar-agar, calcium carbonate, potato or tapioca starch, alginic acid, certain silicates, and sodium carbonate; solution retarding agents such as paraffin; absorption accelerators such as quaternary ammonium compounds; wetting agents such as, for example, cetyl alcohol and glycerol monostearate; absorbents such as kaolin and bentonite clay; and, lubricants such as talc, calcium stearate, magnesium stearate, solid polyethylene glycols, sodium lauryl sulfate, and mixtures thereof.

A skilled person will recognize many suitable variations of the methods to be substituted for or used in addition to those described above and in the claims. It should be understood that the implementation of other variations and modifications of the embodiments of the invention and its various aspects will be apparent to one skilled in the art, and that the invention is not limited by the specific embodiments described herein and in the claims. Therefore, it is contemplated to cover the present embodiments of the invention and any and all modifications, variations, or equivalents that fall within the true spirit and scope of the basic underlying principles disclosed and claimed herein.

The invention having now been fully described, it is exemplified by the following examples and claims which are for illustrative purposes only and are not meant to be further limiting.

Examples

Example 1. Polymers and reagents

Poly(L-lactic acid) (PLLA, M_w ~ 120,000) was purchased from Lactel Biodegradable Polymers (Birmingham, AL). Poly(D,L-lactic-co-glycolic acid) (PLGA) was purchased from Durect, (Cupertino, CA). Dexamethasone (DXM) and HPLC grade acetonitrile were purchased from Sigma-Aldrich (St. Louis, MO). Analytical grade petroleum ether (PE), dichloromethane (DCM), tetrahydrofluran (THF), and glacial acetic acid were purchased from Fisher Scientific (Pittsburgh, PA). PLLA (i.v. = 1.04 dL/g in CHC1₃) and PLGA, 75:25 ester terminated (i.v. = 0.55-0.75 dL/g in CHC1₃) were used in the fabrication of wet spun microfibers. Bovine serum albumin (BSA; Sigma), lysozyme (LZ; Sigma) and bovine zinc insulin (INS; Gibco) were used as model proteins in methods herein. Dichloromethane (Fisher) and petroleum ether (Fisher) were the solvent and nonsolvent used for wet spinning. Sorbitan trioleate (Span 85, HLB 1.8) (Sigma) was used in PLGA microfiber formulations. The micro BCA protein assay reagent kit used to detect protein concentration was used according to the manufacturer (Fisher).

Example 2. Fabrication of dexamethasone-loaded microfibers

PLLA (7.5% w/v), DXM-loaded (1.0%, 2.4%, 4.8% w/w) and unloaded (0% w/w) microfibers were wet spun by phase inversion. PLLA (526.9 ± 0.3 mg) was dissolved in a co- solvent ratio of 6: 1 (v/v) DCM to THF. DXM was added to the co-solvent solution at varying concentrations up to 3.6 mg/mL, near its maximum solubility. The addition of THF was necessary to increase DXM solubility within the spin dope. The polymer /drug solution was loaded into a 5 mL syringe fitted with a 22-gauge spinneret and placed in a syringe pump with a solution flow rate of 0.06 mL/min. Since DCM and THF are miscible with petroleum ether, the immersion of the spinneret into the coagulation bath resulted in continuous microfiber formation and subsequent encapsulation of DXM. A rotating mandrel placed above the spin bath was used to collect microfibers for further analysis. Each formulation in Examples herein was characterized from multiple regions of meter-long microfiber bundles spun at the same time from one spin dope solution. Example 3. Methods of analysis of microfiber morphology

Scanning electron microscopy was used to analyze the surface and cross-sectional morphology of wet spun microfibers. Fiber samples were placed on double-sided carbon tape and coated with a 50-lOOA layer of gold-palladium using a sputter coater (Emitech, Kent, England). Scanning electron microscopy was conducted using a Hitachi S-2700 (Tokyo, Japan) microscope with an accelerating voltage of 8 kV and a working distance of 12 mm. Micrographs were collected using a Quartz PCI digital imaging system. The average microfiber diameter was measured using NIH ImageJ software (Bethesda, MD) from 10 fields of view taken at lOOx magnification. The porosity was measured also from cross-sectioned microfibers and calculated as pore area divided by total cross-sectional area. Example 4. Determination of microfiber drug loading

Drug-loaded and control fiber bundles (5.0 ± 0.3 mg) were dissolved in 5 mL of 6: 1 (v/v) DCM to THF until a clear solution was obtained. The encapsulation efficiency of DXM-loaded microfibers was determined by UV absorbance at 239 nm using quartz cuvettes to minimize background noise at the reading frequency. The percentage of drug encapsulated was calculated as the amount of DXM detected in microfibers relative to the total amount of drug added to the spin dope solution. Two samples from each formulation were assayed in duplicate.

Example 5. Dexamethasone release analysis

As-spun microfibers from each batch (21.0 ± 2.0 mg) were incubated with 2 mL PBS, pH 7.4 in capped microcentrifuge tubes and kept at 37°C. At each time point, 1 mL of the releasate was removed and replaced with fresh PBS. Prior to release analysis, DXM releasates were lyophilized and reconstituted in a mobile phase of 52:48 (v/v) 2 mM acetate buffer (pH 4.8) to acetonitrile. DXM detection was performed using a 3.9 x 150 mm Novapack C-l 8 column with a mobile phase flow rate of 1 mL/min at 240 nm after an average elution time of 2 minutes. Drug concentration was determined by comparing the area under the peak at the expected elution time with a calibration curve constructed from samples of known

concentration. Example 6. Degradation analysis

PLLA microfibers (23.0 ± 2.0 mg) prepared by wet spinning were weighed in microcentrifuge tubes and incubated in 2 mL PBS (pH 7.4). The tubes were capped and placed at 37°C. The duration of the degradation analysis was eight weeks with weekly terminal time points. At each sampling interval, the PBS solution was removed and the supernatant pH measured using a Corning pH meter (Medfield, MA). The remaining microfiber bundles were washed three times in distilled water and lyophilized for 24-48 hours.

Example 7. Differential scanning calorimetry (DSC) and hyper DSC analysis

The thermal properties of microfibers from terminal time points were analyzed for thermal transitions using a DSC-7 (Perkin Elmer) equipped with an Intracooler 2 intercooling system (Perkin Elmer). Samples were subjected to: cooling to -25°C; heating to 250°C at 10°C/min; cooling to -25°C at 10°C/min; and reheating the sample to the upper limit again at the initial rate. Glass transition temperature, melting temperature, and change in enthalpy of the melt were measured from the resulting thermograms. The percent crystallinity (X_c) was also calculated using equation 1 : ΔΗ m,

x l OO

AH P,LLA (1) in which AH_m is the enthalpy of melting of the samples and AH_PLLA (93.7 J/g) is specific heat of melting of a 100% crystalline PLLA as reported in the literature (Fischer EW et al., Kolloid Z Z Polym, 1973;251 :980-990). The dispersion of solid drug particles not solubilized within the polymer matrix was also analyzed using a DSC-8500 (Perkin Elmer) capable of hyper DSC. Samples were subjected to heating from 20°C to 310°C at 200°C/min and compared to the thermogram of free DXM from the manufacturer. Example 8. X-ray diffraction (XRD) analysis

The structural properties of PLLA microfiber formulations were determined using an automated X-ray diffractometer (Siemens Diffraktometer D5000) with a Cu Κα (λ = 1.54A) radiation. The diffraction angles (2Θ) ranged from 6^° to 60^° with sampling intervals of 0.02^° s^"!. Diffraction signal intensity was monitored and processed using DiffracPlus Software (Bruker AXS).

Example 9. Mechanical properties of PLLA micro fibers

Uniaxial tensile testing was used to characterize the mechanical properties of wet spun PLLA microfibers. Tests were performed at ambient temperature, humidity, and pressure using an Instron materials testing machine (Model 4442). Gauge length and elongation rate were defined in accordance to United States Pharmacopeia standards for absorbable sutures ( The United States Pharmacopeia and The National Formulary 201 1. Baltimore, MD: United Book Press, Inc.; 2010 201 1. Baltimore, MD: United Book Press, Inc.; 2010). Microfibers were secured onto paper frames (25 mm x 25 mm) with precut windows to define gauge length and region of loading. After clamping samples into the crossheads of the machine, the edges of each frame were cut leaving microfibers intact. Samples were loaded to failure at a constant elongation rate of 50 mm/min. Load cell measurements and displacement data from crosshead extension were converted into stress-strain data. Ultimate tensile strength, strain at failure, and elastic modulus were calculated from the resulting stress-strain curves. The ultimate tensile strength σ was calculated from the maximum force to failure F divided by the cross sectional area A, using the initial diameter of the fiber obtained from scanning electron micrographs with equation 2:

-1

σ = FA (2) Strain at failure ε was defined as the change in gauge length AL at fiber fracture over the original length of the fiber L₀ using equation 3:

& = ALL₀ ^~X (3) Elastic modulus E was calculated within the linear elastic regime of the stress-strain curve using Hooke's law, whereby strain is linearly proportional to tensile stress using equation 4:

E =σε-^! (4)

Example 10. Statistical analysis

One-way analysis of variance (ANOVA) tests were performed on mechanical data for wet spun microfibers using SPSS v.19 statistical software (Chicago, IL). Post hoc analyses were carried out using the Tukey multiple comparisons test. A p value of less than 0.05 was considered to be significant.

Example 11. Microfiber extrusion and morphology

PLLA spin dope solutions extruded into the petroleum ether coagulation bath underwent phase inversion within centimeters of the spinneret, forming an opaque white microfilament. Residence time was 5-10 seconds before microfibers were collected around a rotating mandrel placed above the spin bath. Samples were strong and spun easily, indicating quite fast and effective curing within this short time span. Representative scanning electron micrographs of microfibers are shown in FIG. 1 panels A and B. Wet spinning conditions produced 64.3 ± 7.0 μηι diameter microfibers with skin-core structures. Skin-core structure is common to wet spinning, and is caused by rapid surface coagulation leading to the entrapment of solvent and nonsolvent within the precipitating microfilament (Rissanen M et al., J Appl Polym Sci, 2008; 1 10:2399-2404). Porous structure is formed once solvent and nonsolvent are evaporated after microfilament solidification. The solvents involved in the formation of the fiber play a major role in further crystallizing the polymer by solvent-induced crystallization (SINC). DXM- loaded and control microfibers were observed to have round cross-sectional geometry and relatively porous morphology with similar interconnectivity.

The average percent porosity of microfibers was observed to be 17.0 ± 1.3 with an average pore size area of 1.26 ± 0.06 μιη² as determined from the micrographs of cross- sectioned microfibers. Microfibers displayed micron-rough surface topography with spherical protrusions similar to studies by Nelson et al. (Nelson KD et al., Tissue Eng, 2003 ;9: 1323- 1330). Spherical structures were observed, such as spherulites (FIG. 1) forming as the polymer is cured (Cohen Y et al, Acs Sym Ser, 1987;350: 181 -98; Sukitpaneenit P et al., J Membrane Sci, 2009;340: 192-205; Teuji H et al., Macromolecules, 1992;25:2940-2946). Maltese cross patterns were observed under cross-polarized light (FIG. 8).

Incubation time was observed to have no significant effect on microfiber shape and cross-sectional morphology (FIG. 1 panels C and D). Diameters did not change more than 5.0 μιη after incubation in PBS for eight weeks. The average microfiber diameters of 0% (control), and 1.0%, 2.4%, and 4.8% (w/w) DXM-loaded samples throughout the degradation analysis were observed to be 62.8 ± 6.8 μηι, 65.7 ± 5.8 urn, 65.6 ± 5.8 μηι, and 62.9 ± 8.3 μηι, respectively. An occasionally large pore was observed among sample cross sections. The overall porosity of fibers did not change with time. At eight weeks of incubation, microfibers had an average percent porosity of 18.3 ± 3.1 and an average pore size area of 1.24 ± 0.09 μιη². These data indicate that little PLLA microfiber degradation had occurred during the incubation period.

Example 12. Dexamethasone encapsulation and in vitro release kinetics

To determine the encapsulation efficiency and drug distribution within wet spun microfibers, two separate regions of microfiber bundles collected from the coagulation bath were assayed. DXM 1.0-4.8%) (w/w) was encapsulated with relatively high efficiency (76-93%) as shown in Table 1. DXM and PLLA were dissolved in 6: 1 (v/v) DCM to THF before precipitation in the petroleum ether bath. The coagulation rate of fibers was nearly

instantaneous, resulting in little overall drug loss during polymer solidification. DXM-loaded microfibers at 4.8% (w/w) were observed to have slightly decreased encapsulation efficiency in comparison to 1.0% and 2.4% (w/w) formulations. The observed decrease with 4.8% (w/w) drug loading was due to obtaining the limit for drug loading of 7.5% (w/v) PLLA microfibers; drug loadings greater than 4.8% (w/w) resulted in significant drug loss as shown with 10% (w/w) theoretical DXM loading in Table 1. For these reasons, formulations with 1.0%, 2.4% and 4.8% (w/w) theoretical loading were used in Examples herein. No significant variations in drug loading were observed among different sampling regions of meter-long microfibers bundles indicating that the drug was homogeneously distributed.

Table 1. Encapsulation efficiency of microfibers prepared by wet spinning. Data are represented as mean ± S.D.

DXM loading of dope DXM loading of wet spun DXM encapsulation

solution (%w/w) microfiber (%w/w) efficiency (%)

1.0 0.9 89.0 ± 0.036

2.4 2.2 92.9 ± 0.084

4.8 3.6 75.8 ± 0.002

10.0 1.0 9.9 ± 0.010 Release profiles of DXM-loaded fibers are shown in FIG. 2 panel A. Each fiber formulation exhibited little burst, followed by linear cumulative percent DXM release starting at day 3 and continuing to eight-week incubation (R2 = 0.92-1.0). Formulations with 0.9% and 2.2% (w/w) actual DXM loading released less than 1% and 2% of total drug encapsulated during the first three days. In contrast, 3.6% (w/w) DXM-loaded fibers released 5% of total drug encapsulated during the first three days and maintained a faster rate of drug release throughout the course of the analysis. Greater initial release of DXM from 3.6% (w/w) actual drug-loaded fibers was due to drug located near the surface of microfibers. Increased drug loading led to an overall increase in drug release, with lower loadings displaying similar release rates as a function of time. These data indicate that at lower loadings, the drug was well encapsulated within the polymer matrix with little drug near the surfaces of microfibers. The cumulative percent DXM released after eight weeks was 5.54 ± 0.12 for 0.9% (w/w) actual drug loading, 4.62 ± 1.08 for 2.2% (w/w) actual drug loading, and 27.82 ± 5.13 for 3.6% (w/w) actual drug loading.

An absence of morphological changes was observed by scanning electron microscopy. These data indicated that drug release was not predominately controlled by polymer degradation, and rather by dissolution and diffusion of drug into the supernatant. At low loadings, drug release was associated with dissolution of the drug, and at higher loadings, drug release was associated with a combination of dissolution and diffusion. Since the 3.6% (w/w) actual drug- loaded microfibers released 202.5 μg after eight weeks, nearly ten times the amount of drug relative to the 2.2% (w/w) formulation that released 20.6 μg, a significant portion of drug release was due to diffusion. Mack et al. have also studied release of DXM from microfibers (Crow BB et al., Biopolymers, 2006;81 :419-427; Mack BC et al, J Control Release, 2009;139:205-211). At eight weeks, wet spun PLGA (50:50) microfibers released 60-90% total encapsulated drug prim arily due to polymer degradation. According to the manufacturer, the degradation rate of the raw PLLA used in Examples herein is greater than 24 months. Therefore, the release behavior from microfiber delivery system observed herein over the course of eight-week incubation was predominantly due to the hydrophobic nature of DXM.

The controlled, sustained release of dexamethasone constitutes a clinically important therapeutic modality for biomaterial-driven inflammation. Local, targeted delivery of dexamethasone over extended periods of time (months) reduces immune responses to implanted biomedical devices and polymeric biomaterials (Hickey T et al., J Biomed Mater Res,

2002;61 : 180-187; Patil SD et al., Diabetes Technol Ther, 2004;6:887-897; Dang TT et al, Biomaterials, 201 1 ;32:4464-4470; Selvam S et al, Biomaterials, 201 1 ;32:7785-7792). The effective therapeutic ranges of drug evaluated in previous studies ranged from 0.17 to 30 g/day, depending on the size and material of the implant. The average daily release rates after initial burst for 20 mg microfiber bundles from compositions and methods observed herein were approximately 1 -2 μg/day for 0.9 and 2.2% actual DXM loading, and 15 μg/day for 3.6% actual DXM loading. These data indicate an ability to tune DXM release rates within the therapeutic ranges of drug treatment by combining multiple formulations into multifilament yarns and by altering scaffold architecture (Leung V et al., Polym Advan Technol, 201 1 ;22:350-365).

For example, weft-knitted scaffolds use more yarn material in the assembly of macrometric structures and thus are capable of delivering a greater dose of therapeutics than warp-knitted scaffolds of similar dimensions.

Example 13. Supernatant pH characteristics of DXM-loaded PLLA microfibers

Diffusion of PLLA oligomer was analyzed by measuring the supernatant pH over degradation time. Degradation of PLLA (Mathiowitz E., Encyclopedia of controlled drug delivery. New York: Wiley; 1999) yields L(+)-lactic acid, a naturally occurring stereoisomer of lactic acid. The lactic acid monomers from hydrolytic de-esterfication of PLLA enter the carboxylic acid cycle and are excreted as water and carbon dioxide. For this reason PLLA is widely used in drug encapsulation.

The supernatant pH of terminal time points was determined to compare the relative changes of DXM-loaded microfibers in comparison to control microfibers with no drug.

Formulations displayed similar decrease in pH (FIG. 2 panel B). At eight-week incubation, 0% (control), 1.0%, 2.4% and 4.8% (w/w) formulations decreased in pH by 0.95, 1.25, 1.56 and 1.37, respectively. Inclusion of DXM potentially increases the hydrophobicity of microfibers, resulting in little pH changes as a function of time. The PBS was not changed and each time point was terminal for this Example only. Crow et al. measured supernatant pH changes as a function of bovine serum albumin (BSA) release from wet spun PLLA microfibers (Crow BB et al., Tissue Eng, 2005;1 1 : 1077-1084) and observed that at 15-week incubation, control microfibers decreased in pH by 1.70, with buffer changes occurring at each time point. Data herein indicated that DXM release from microfiber was not mediated by PLLA degradation.

Example 14. Thermal analysis of DXM-loaded PLLA microfibers

The morphology of blank and degraded microfibers was analyzed using DSC.

Representative DSC traces of control microfibers after fabrication and throughout degradation are presented in FIG. 3 panel A. Thermal transitions calculated from the first heating scan are listed in Table 2. It was observed that relative PLLA crystallinity, as a result of wet spinning, was substantially increased. Extrusion of PLLA alone increased relative crystallinity by 17% in comparison to raw polymer from the manufacturer. Drug-loaded microfibers also increased in relative crystallinity by 13 -16% after wet spinning. Several factors contributed to this effect, including the solvents, nonsolvent, and residence time during fabrication, leading to SINC (FIG. 4 panel B) (Neogi P., Diffusion in polymers. New York: Marcel Dekker; 1996). Many polymers undergo SINC, and as a polymer (semi-crystalline) is exposed to a solvent between the T_g and the T_m, crystallization becomes kinetically favorable. The choice of the solvents and the solubility parameters of solvent [5DCM ⁼ 20.2 megapascals (MPa ), 5THF ⁼ 18.6 MPa ] and nonsolvent (5PE =14.8 MPa^1/2) in our system affected the rate of solvent removal, solidification, and eventually the degree of crystallinity of PLLA (5PLLA =20.7 MPa^{1 2}) microfilaments. In wet spinning, an initially homogeneous polymer solution is extruded into a coagulation bath that induces phase inversion by counter-diffusion of solvent and nonsolvent. Phase separation of the polymer solution is then initiated and the liquid spin dope stream begins to precipitate into a solid microfilament. In general, it is possible to associate the counter-diffusion of solvent and nonsolvent with the thermodynamic stability of extruded spin dope solutions (Sukitpaneenit P et al., J Membrane Sci, 2009;340: 192-205); the precipitation of polymeric solutions is typically faster with a higher diffusion of solvent in nonsolvent. Wet spun microfibers with similar

1/2 residence time prepared by issanen M et al by phase inverting PLDLA (5PLDLA = 23.3 MPa ) dissolved in DCM into an ethanol (EtOH) spin bath (δ_ΒΟΗ = 26.6 MPa^{1 2}) resulted in a 29% decrease in crystallinity (Rissanen M et al., J Appl Polym Sci, 2009;113:2683-2692). The difference in the solubility parameter between the solvent /nonsolvent, in the last Example, was about 6 units, which indicating that ethanol is a better nonsolvent for the microfiber than petroleum ether. Thus, slower precipitation is optimal for SINC.

No significant differences in glass transition temperature after fabrication were observed between the 1.0% and 2.4% (w/w) DXM-loaded samples in comparison to control microfibers (Table 2). Surprisingly, 4.8% (w/w) DXM-loaded samples showed a slight decrease in glass transition temperature. This indicates that DXM did not act as plasticizer and interacts with the polymer matrix, or that too little drug was present to detect changes in glass transition temperatures. Since there appeared to be little DXM interaction with the amorphous regions of the polymer, as judged by the minor changes in glass transition, Hyper DSC was used to analyze the dispersion of solid drug within wet spun microfibers. Conventional DSC techniques limit the sensitivity to measure the drug melting temperature (Gramaglia D et al., Int J Pharm,

2005;301 :1-5). Slow heating scan rates of 10°C/min allot sufficient time for molecules within the crystal lattice to respond to thermal transitions and further solubilize within the polymer matrix. Hyper DSC uses scanning rates of >100°C/min to inhibit further solubilization of drug within the matrix. Therefore, the fraction of drug solubilized within the matrix does not contribute to the melting endotherm associated with the dispersed drug fraction. Hyper DSC thermograms revealed no melting point at 300°C for DXM (FIG. 3 panel B), indicating the drug was either amorphous or formed a solid solution within PLLA microfibers and precipitated in an amorphous state. The results of Hyper DSC thermograms indicate that the drug was dispersed at the molecular level since the glass transition of the highest drug loaded fiber, 4.8% was lower than the blank spun fibers. X D analysis (Example 15) also showed no crystalline content of DXM in the nanoparticles. Table 2. Summary of thermal properties and crystallinity from the first DSC heating scan of degraded microfibers

Degradation

T_g(°C) T_m (°C)

time (weeks) ¾ (%)

Control (0% DXM)

0 57.4 174.2 60.5 64.6

2 61.6 174.3 60.6 64.7

4 63.0 174.2 60.6 64.6

6 63.1 174.2 60.9 65.0

8 64.3 173.5 59.2 63.2

1% DXM

0 59.5 173.7 59.2 63.2

2 55.9 173.4 59.2 63.1

4 63.3 174.7 56.0 59.8

6 63.8 174.0 61.1 65.2

8 62.2 173.9 61.4 65.5

2.4% DXM

0 59.7 173.4 56.4 60.1

2 62.9 173.3 59.0 63.0

4 62.6 174.2 57.8 61.7

6 62.4 173.9 60.8 64.9

8 64.4 173.5 59.8 63.8

4.8% DXM

0 54.3 173.7 57.0 60.8

2 59.4 174.0 57.5 61.4

4 62.2 174.0 58.9 62.8

6 63.2 173.4 58.4 62.4

8 63.6 174.0 57.9 61.8

Unprocessed PLLA

n/a 52.6 174.7 44.5 47.4

T_g, glass transition temperature; T_m, melting temperature; AH_m, endothermic enthalpy of the melting peak; X_c, crystallinity degree.

The polymer crystallinity and glass transition temperatures of degraded samples taken from terminal time points were also evaluated using DSC. If major degradation had occured, the amorphous regions would mostly be expected to disappear, leaving fragmented samples with very high crystallinity. Thus, an increase in crystallinity would be expected. However, DXM- loaded samples were observed to have minimal changes in crystallinity as a function of time, signifying minor polymer degradation. The glass transition temperature of the microfibers slightly increased with incubation. Polymers which are annealed display an increase in glass transition which is related to a reduction in free volume and possibly segmental rearrangement. The glass transition temperature for the second heating scan of the raw polymer was observed to be 47.6°C. Since the glass transition temperature of the second heating scan is close to the incubation temperature (37°C), the mobility of the polymer chains had increased with incubation time. An increase in chain mobility would result in the formation of crystalline structures (Zong X et al., Biomacromolecules, 2003;4:416-423). Constrained amorphous chains from thermally induced crystallization resulted in the increase in glass transition temperature with incubation. Secondary DXM-polymer interactions already existing from fabrication would have further increased as some of the polymer degraded. Even minor PLLA degradation during incubation resulted in free carboxylic and hydroxylic groups that are capable of forming secondary interactions with the drug (Fig. 4 panels B and C), contributing to amorphous chain restructuring and a reduction in free volume. Additionally, the specific interactions between PLLA and DXM contributed to the slow release of drug from this delivery system. Example 15. X-ray diffraction (XRD) analysis of DXM- loaded PLLA microfibers

XRD was used to further investigate the physical state of the drug in wet spun microfibers. No crystalline DXM was detected in the four microfiber formulations namely control (0%), 1.0%, 2.4% and 4.8% w/w DXM-loaded PLLA microfibers (FIG. 5). The absence of the specific diffraction of DXM in the x-ray diffractogram of drug-loaded PLLA microfibers demonstrated that DXM was present as either an amorphous drug or as solid solution within the amorphous regions of the polymer (molecular dispersion). Evaluation of the formulations herein showed that control microfibers without drug had a less ordered state than microfibers loaded with drug.

As drug loading was increased the crystalline order increased and a new crystalline peak appeared in the diffraction pattern as shown with 1.0, 2.4, and 4.8% drug loading in FIG. 5.

Since microfibers had short residence time (5-10 seconds) in the nonsolvent, residual solvent was likely present within the precipitating filament after it was collected from the spin bath.

SINC is associated with the increase in the AH_m of the blank (control) fibers compared to pure polymer. The change in the PLLA XRD morphology with the drug-loaded samples resulted from the solvent-drug interaction during SINC that keeps the glass transition lower for a longer time, thus allowing for further crystallization to take place after the precipitation has started, associated with the new peaks in FIG. 5 for the drug-loaded formulations. The control fibers left in the nonsolvent bath for a longer residence time revealed the same third peak (FIG. 9) and the higher degree of crystallinity observed in drug-loaded microfiber formulations described herein.

Example 16. Mechanical properties of DXM- loaded PLLA micro fibers

To characterize the mechanical properties of wet spun microfibers, samples were loaded in uniaxial tension until failure. The results of mechanical analysis are shown in FIG. 6 and summarized in Table 3. Stress-strain curves for microfibers from each formulation displayed large plastic deformation until failure. Formulations were observed to have similar elastic moduli (FIG. 6 panel A). Differences in ultimate tensile strength (UTS) and the total elongation before tensile failure (strain at failure) were observed between drug-loaded samples compared to control samples without drug. Microfibers with the highest drug loading, 4.8% (w/w) were observed to have the greatest tensile strength (15.6 ± 0.5 MPa) of drug-loaded formulations and were significantly stronger (p < 0.01 ) than control microfibers (FIG. 6 panel B). Drug encapsulation at low loadings, 1.0% and 2.4% (w/w), resulted in similar tensile strengths in comparison to control microfibers. Strain at failure for formulations containing drug was significantly different, with increased drug encapsulation resulting in increased plastic deformation until fracture (FIG. 6 panel C). The highest drug loading, 4.8% (w/w) displayed similar elongation until failure as control microfibers containing no drug.

Table 3. Summary of the mechanical properties of wet spun fibers, mean ± SEM

DXM loading of Tensile strength Elastic modulus Strain at Failure dope solution (%) UTS(MPa) (MPa) mm/mm

0 12.9 ± 0.8 195 ± 25.5 1.78 ± 0.13

1.0 14.0 ± 0.6 215 ± 26.1 1.09 ± 0.09

2.4 1 1.7 ± 0.4 204 ± 20.0 1.29 ± 0.10

4.8 15.6 ± 0.5 213 ± 24.8 1 .70 ± 0.10 Necking during the macroscopic plastic deformation of wet spun microfibers was observed to grow stronger due to the alignment of polymer chains with increasing stress. In semi-crystalline polymers, the amorphous regions of the polymer required less force to deform in comparison to the crystalline regions of the polymer. Increased strength and strain with the 4.8% (w/w) formulation was associated with dexamethasone reinforcing the amorphous regions of the polymer. The process of wet spinning resulted in a marked increase in crystallinity as determined by DSC. Without being limited by any particular theory or mechanism of action, with closely packed and parallel polymer chains, i.e. polymers with increased crystallinity, mechanical strength increased in the presence of drug due to enhanced secondary bonding. During incubation, dexamethasone molecules (392.8 Da) formed secondary interactions within the amorphous phase of the polymer, creating a reinforced composite material (FIG. 4 panel C). With small amounts of encapsulated dexamethasone, the strain was observed to decrease until sufficient drug was encapsulated to exhibit similar strain at failure as control microfibers.

Without being limited by any particular theory of mechanism of action it was envisioned that drug particles acted as material defects at low encapsulation loading. Increasing DXM loading increased drug-drug and drug-matrix interactions. Therefore, mechanical properties were influenced by the secondary interactions, namely hydrogen bonds between the polymer lattice and the drug particles.

To further elucidate the effect of drug loading and release on microfibers, mechanical properties were characterized throughout the degradation analysis. Tensile loss profiles of each formulation were generated from stress-strain profiles of terminal time points (FIG. 7 panel A). Microfibers with the highest drug loading, 4.8% (w/w), maintained 97.3% of initial tensile strength and were statistically stronger (p < 0.05) after eight weeks in vitro (FIG. 7 panel B). Control, 1.0% and 2.4% (w/w) microfibers were observed to have decreased tensile strength in the first week, and then to have maintained strength. At the end of eight weeks, the breaking strength retention for control, 1.0%, and 2.4% (w/w) was observed to be 80.9%, 80.0%, and 81.3%, respectively. Formulations displayed a decrease in plastic deformation and strain at failure as a function of time (FIG. 7 panel C) independent of drug content, see measurement of the specific heats reported in Table 2. The crystallinity of wet spun microfibers increased as a function of time, causing microfiber embrittlement. Elastic moduli of samples and resistance to deformation under load were conserved as a function of time (FIG. 7 panel D). Dexamethasone drug particles at 4.8% (w/w) were observed to contribute to mechanical strength by secondary bonding adding to physical reinforcement. DXM also was observed to preserve the structure of microfibers as a function of time due to drug stabilizing the lattice, and decreasing the permeation rate due to hydrophobic drug content.

Studies with melt spun PLLA microfibers encapsulating a small anti-inflammatory drug, curcumin (368.4 Da), have also shown preservation of mechanical properties with drug loading (Su SH et al., J Biomat Sci-Polym, E 2005; 16:353-370). Curcumin impregnation at 10% (w/w) increased fiber tensile strength at failure for periods up to 36 days in vitro in comparison to control microfibers with no drug. Although melt spun microfibers produced higher tensile strengths than that of wet spun microfibers in Examples herein, curcumin loaded microfibers maintained only 57% of their initial tensile strength at 25-day incubation. The processing conditions of wet spun microfibers fabricated by Mack et al. also possessed higher initial tensile strength than those used in Examples herein (Mack BC et al., J Control Release, 2009; 139:205- 21 1). However, wet spun microfibers encapsulating levoxithan and dexamethasone lost nearly all of their tensile strength after only 7-day incubation. Data herein show that dexamethasone incorporation increased mechanical integrity as a function of increased microfiber crystallinity. Melt spinning is known to increase crystallinity due to the manner by which polymer chains group themselves in forming a microfiber and therefore curcumin impregnation helped maintain mechanical properties. Conversely, wet spun microfibers fabricated using DMSO and water as solvent and nonsolvent reported no increase in crystallinity and therefore dexamethasone incorporation did not help maintain mechanical strength. Addition of small hydrophobic drugs (<400 Da) increased mechanical integrity of filaments as crystallinity is increased as a result of fabrication. As the proximity of amorphous and crystalline regions increases, so does the possibility for strong secondary interactions between drug and polymer. Example 17. Microfiber fabrication for encapsulating proteins

Spin dope solutions were prepared using a modified cryogenic emulsion technique (Mathiowitz E et al., Nature, 1997;386:410-414). PLGA and PLLA microfibers were loaded with a protein: bovine zinc insulin (INS, 5.8 kDa), lysozyme (LZ, 14.3 kDa), or bovine serum albumin (BSA, 66.0 kDa). Protein in ultra-pure water 0.5 mL (20 mg mL^"1) was added to 10 mL of polymer in DCM (50 mg mL^"1), yielding an aqueous to organic phase ratio of 1 :20. This two- phase system was vortexed for 60 s to create a meta-stable emulsion. The emulsion was frozen in liquid nitrogen, creating frozen protein droplets dispersed in frozen dichloromethane/polymer solid solution. The frozen emulsion was lyophilized for 48 h at -100°C. This process resulted in proteins particles imbedded in a matrix of a polymer of less than 2 μηι. The W/O (water/oil) emulsion micronization method showed that most proteins have a solid size <2 μιη at the end of the process regardless of the nature of the protein (reference here is to the physical size of the particle, not the molecular weight of the protein) (Mathiowitz E et al., United States Patent. USA: Brown University Research Foundation; 2006).

The dried polymer and protein product was reconstituted in 2.5 mL DCM at a concentration of 200 mg mL^"1 and placed into a gas-tight glass syringe fitted with a 22-gauge spinneret. A syringe pump was used to extrude the spin dope solution (0.02-0.06 mL min^"1) into petroleum ether at a solvent to nonsolvent ratio of 1 :400, which resulted in the continuous formation of PLGA and PLLA monofilaments. PLGA formulations required 50 of Span 85 to prevent gelation around the spinneret during extrusion. Extruded microfibers were collected from the coagulation bath after the spin dope was extruded (aboutl .5 h residence time). Blank microfibers were also fabricated and used as controls. Duplicate batches of each formulation were made and analyzed to ensure the reproducibility and uniformity of release profiles.

Example 18. Analysis of protein loaded microfibers by Scanning electron microscopy

Scanning electron microscopy was used to analyze cross-sectional and surface

morphology of wet spun microfibers. Lyophilized microfibers were mounted on adhesive metal stubs and sputter-coated with a 50-100 A layer of gold-palladium (Emitech). Samples were viewed with a Hitachi S-2700 scanning electron microscope using an accelerating voltage of 8 kV. Micrographs were taken using a Quartz PCI digital imaging system. To determine average cross-sectional area, microfibers were arranged into bundles onto paraffin film and rolled laterally into cylindrical tubes and orthogonally cut into thin discs. Five fields of view (15-22 filaments) of cross-sectioned microfibers were captured at 350x magnification and analyzed using ImageJ software (NIH).

Formulations were observed to have a phase inversion of about 2-4 cm from the spinneret tip, forming a continuous solid white monofilament. Microfibers were observed to have consistent size, and similar cross-sectional area, porosity, and porous interconnectivity. Blank and protein-loaded PLGA microfibers were observed to have lobed 'dog-bone' shape with an average height of 46 μηι and width of 105 μιη (FIG. 10). Blank and protein-loaded PLLA formulations displayed lobed 'kidney' shape with an average height of 50 μηι and width of 102 nm (FIG. 1 1).

Example 19. Thermal analysis of protein loaded microfibers by DSC

DSC was used to analyze the thermal properties of PLGA and PLLA microfibers after fabrication (day 0) and the incubation period (day 63). DSC measurements of blank and protein- loaded microfibers after fabrication (day 0) and incubation (day 63) were conducted using a DSC-7 (Perkin Elmer) equipped with an Intracooler 2 intercooling system. Samples were subjected to: cooling to -25°C, heating to 250°C at 10°C min^"1, cooling to -25°C at 10°C min and then reheating to the upper limit again at the initial rate. The crystallinity (X_c) of PLLA microfibers was calculated using X = s^. x 100

93.7 (5)

in which 93.7 J g^"1 is the specific heat of melting of a 100% crystalline PLLA (Fischer EW et al.,

Kolloid Z Z Polym, 1973;251 :980-990).

Representative DSC curves of control PLGA and PLLA microfibers after fabrication and after 63-day incubation are presented in FIG. 12. Table 4 summarizes the thermal transitions of wet spun microfibers from the first DSC heating scan. The glass transition temperature (T_g) of blank PLGA microfibers was 40.7°C. The T_g of microfibers slightly increased (1.3-3.7°C) with protein loading. After 63-day incubation, the PLGA formulations, including blank microfibers exhibited an increase in J^ in the range of 2.4-5.0°C.

Table 4. Summary of the thermal properties from an initial DSC heating scan of wet spun microfibers upon fabrication (0 days) and after incubation (63 days). Data are expressed as mean ± S.D. (n = 2).

Microflber Time T_g (°C) T_m (°C) AH_m (J g-') Xc (%)

PLGA

Blank t = 0 40.7 ± 0.5 - - - t = 63 45.3 ± 1.1 - - -

INS t = 0 44.4 ± 3.8 - - - t = 63 46.8 ± 0.6 - - -

LZ t = 0 42.4 ± 1.9 - - - t = 63 47.4 ± 0.7 - - -

BSA t = 0 42.0 ± 1.5 - - - t = 63 46.3 ± 0.4 - - -

PLLA

Blank t = 0 42.1 ± 2.2 173.0 ± 0.1 24.6 ± 2.4 26.2 ± 2.6 t = 63 58.2 ± 0.7 173.3 ± 1.5 28.6 ± 1.0 30.5 ± 1.1

INS t = 0 40.8 ± 0.4 173.0 ± 1.5 27.5 ± 0.4 29.4 ± 0.4 t = 63 59.1 ± 0.5 172.0 ± 0.6 27.5 ± 0.3 29.3 ± 0.3

LZ t = 0 42.4 ± 2.5 173.0 ± 1.4 27.0 ± 0.5 28.8 ± 0.5 t = 63 58.9 ± 0.7 172.0 ± 0.0 28.4 ± 1.7 30.3 ± 1.8

BSA t = 0 40.7 ± 0.6 173.3 ± 1.3 24.3 ± 0.6 25.9 ± 0.6 t = = 63 59.2 ± 0.5 171.3 ± 0.6 28.4 ± 0.3 30.3 ± 0.3 Control (blank) PLLA microfibers were observed to have a T_g of 42.1 °C, with a T_m of

173.0°C, and relative crystallinity of 26.2%. The T_g of INS- and BSA-loaded microfibers was observed to be slightly decreased by 1.3°C and 1.4°C compared to the control; LZ microfibers were similar to the control. A minor increase in relative crystallinity (2.6-3.2%) was observed with INS and LZ-loaded PLLA microfibers compared to blank PLLA microfibers. BSA-loaded microfibers were observed to have the lowest relative crystallinity of 25.9% and T_g of 40.7°C. In contrast to PLGA, PLLA (63 day) microfibers were observed to have substantially increased T_g of about 16.1 to 18.5°C. PLLA microfibers also were observed to have a slight increase in percent relative crystallinity; blank microfibers and BSA-loaded microfibers were observed to have the greatest increase in percent relative crystallinity of 4.3% and 4.4%, respectively. The T_m for microfibers formulated with PLLA was maintained at an unchanged level during the incubation period. Example 20. X-ray diffraction (XRD) analysis of protein loaded microfibers

XRD was used to determine the crystalline structure of PLLA microfibers. The structural properties of PLLA formulations at the initial time point and at 63 days of incubation were also determined using an automated X-ray diffractometer (Siemens Diffraktometer D5000) with a Cu α (λ = 1.54A) radiation. Diffraction was measured at diffraction angles (2Θ) between 6 and 60 with sampling intervals of 0.02 s^"1. Diffraction signal intensity was monitored and processed using DiffracPlus Software (Bruker AXS).

X-ray scattering patterns of control (blank) and protein-loaded microfibers before (day 0) and after incubation (63 days) showed two small peaks (13.0° and 17.6°) and one large peak (24.4°) for PLLA microfibers (FIG. 3 panel A). After incubation in PBS at 37°C for 63 days, the diffraction peak at 24.4° was no longer observed and a new diffraction peak between 16-17° was present, which was significantly enhanced for protein-loaded samples (FIG. 13 panel B). Example 21. Mechanical testing of protein loaded microfibers

To determine the effect of protein loading and molecular weight on the mechanical properties of PLGA and PLLA wet spun microfibers, samples were loaded under uniaxial tension until failure. Uniaxial tensile tests were conducted using a materials testing system (Instron Model 4442) in accordance to the United States Pharmacopeia absorbable suture testing standard (United States Pharmacopeia and National Formulary (USP 34-NF 29, Rockville, MD: United States Pharmacopeial Convention; 2010). Microfibers and yarns were secured to a paper frame (25 mm x 25 mm) and loaded into the crosshead clamps of the machine. Prior to loading, the sides of the paper template were cut leaving the sample intact. An elongation rate of 50 mm min^"1 was applied until failure. The resulting load-displacement data collected by the digital acquisition system was converted to stress-strain data to calculate the ultimate tensile strength (UTS), percent strain to failure, and elastic modulus. The mechanical properties of microfibers were compared in SPSS v.19 (Chicago, 1L) using ANOVA. Analyses for multiple comparisons were carried out using the Tukey multiple comparisons test. A p value of less than 0.05 was considered statistically significant.

Since microfibers were non-circular, the average cross-sectional areas were calculated from scanning electron microscopy images for mechanical analyses. The mechanical strength of blank PLLA microfibers was observed to be greater than control (blank) PLGA microfibers (FIG. 14). Blank PLLA microfibers were observed to have an average load to failure of 176.4 ± 3.8 millinewtons (mN) and ultimate tensile strength of 43.2 ± 0.9 MPa, compared to

blank/control PLGA microfibers that were observed to have an average load to failure of 1 18.6 ± 14.0 mN and ultimate tensile strength of 32.7 ± 3.8 MPa (FIG. 14 panels A and B). Significant variations in the strain to failure and elastic modulus were also observed among blank formulations. PLGA microfibers had a strain to failure of 12 ± 6% in comparison to PLLA microfibers with a strain to failure of 3 ± 1% (Fig. 14 panel C). PLGA microfibers had an average elastic modulus of 569.7 ± 61.7 MPa whereas PLLA microfibers had an average elastic modulus of 822.4 ± 93.3 MPa (Fig. 14 panel D).

Protein encapsulation was observed to surprisingly alter the material properties of both PLGA and PLLA microfibers. Protein-loaded microfibers within the same polymer type had significantly lower (p < 0.05) load to failure and ultimate tensile strength compared to blank controls (Fig. 14 panels A and B). Among protein-loaded formulations, BSA-loaded PLGA and PLLA microfibers possessed significantly lower (p < 0.05) load to failure and ultimate tensile strength compared to INS and LZ-loaded formulations. However, no statistical differences in load to failure and ultimate tensile strength were observed between INS and LZ-loaded PLGA formulations.

Significant differences in strain at failure were also observed within PLGA and PLLA formulations (FIG. 14 panel C). The failure extensions of protein-loaded PLGA microfibers were statistically lower than blank PLGA microfibers (p = 0.000). LZ and BSA-loaded PLLA microfibers also had statistically lower elongation until failure compared to blank PLLA microfibers. INS-loading did not significantly decrease the ductility of PLLA microfibers (p = 0.659). An exponential relationship was found between parameter of the strain to failure and molecular weight of the protein that was encapsulated in microfibers for each of PLGA (R² = 0.9880) and PLLA (R² = 0.9800). No statistical differences in elastic moduli were observed among PLGA formulations (Fig. 14 panel D). However, BSA-loaded PLLA microfibers were observed to have significantly lower elastic modulus in comparison to blank, INS-loaded and LZ-loaded formulations.

Example 22. Analysis of protein release from microfiber formulations

To determine protein release kinetics, triplicate samples (10 mg) from each batch were suspended in 1.5 mL of PBS (pH 7.4) in capped microcentrifuge tubes and incubated at 37°C. At each time point, 1 mL of the releasate was removed and replaced with fresh PBS. The pH of the release buffer was 7.2-7.4 throughout the incubation period. The protein concentration of releasates and corresponding protein standards in PBS (1.6-25 μg mL^"1) were determined using a micro BCA assay kit. Samples and corresponding standards were read at 562 nm. In vitro anayses were performed in triplicate from duplicate batches of each formulation. The release profiles are expressed as the cumulative percent release from each formulation, normalized to the amount of encapsulated protein. Microfibers from each formulation were immersed in 0.1 N NaOH (Sigma St. Louis, MO) and incubated at 37°C until fully dissolved. The encapsulation efficiency was calculated as the actual amount of protein detected with the micro BCA assay relative to the theoretical amount of protein added to spin dope solutions prior to wet spinning.

The data obtained from in vitro protein release was fitted to the Korsmeyer-Peppas kinetic model to determine the mechanism of release. Korsmeyer et al. derived a simple relationship to describe diffusion based drug release from a polymeric system given by

-^^L = kt" (6)

M,

in which—— is a fraction of drug released at time /, k is the release rate constant, and n is the

Μ_Λ

release exponent (Korsmeyer RW et al, Macromolecular and modeling aspects of swelling controlled system. In: Roseman TJ, Mansdorf SZ, editors. Controlled release delivery systems. New York: M. Dekker; 1983, 77-101). In this model, the n value is used to characterize the release mechanism. For a cylindrical matrix, n < 0.45 corresponds to Fickian diffusion, 0.45 < n < 0.89 to non-Fickian transport, n = 0.89 to case II (relaxational) transport, and n > 0.89 to super case II transport. Data were plotted as log cumulative percentage drug release as a function of log time for the first 60% of drug released.

Three phases of release from microfiber formulations were observed (Table 5). During the first phase, protein release was fastest for INS, followed by LZ and BSA for PLGA formulations (FIG. 15 panel A). Similar patterns were observed for PLLA formulations, but at significantly lower cumulative release percentages (FIG. 15 panel B). The duration of the first phase was characterized by a typical diffusion profile and was found to be dependent on protein molecular weight. For PLGA microfibers, this phase lasted for 14 days for INS, and 21 days for LZ and BSA. The first phase of INS, LZ and BSA release from PLLA fibers was 14 days. A second phase, characterized by slow, linear release was also observed for the PLGA and PLLA formulations. The duration of this phase for INS was 28 days and 21 days for PLGA and PLLA formulations, respectively. LZ and BSA PLGA formulations were observed to have a second phase of 42 days; PLLA formulations were observed to have a second phase of 49 days. A third phase with additional protein release was observed with INS formulations, which lasted 21 days for PLGA and 28 days for PLLA formulations.

To determine the mechanism of protein release from microfibers, in vitro release data were fitted to the Korsmeyer-Peppas kinetic model. The regression coefficient (R²) and release constant (k) values from the release data for each formulation were calculated as shown in Table

6. The regression coefficients for Korsmeyer-Peppas plots were observed to be in a range from Table 5. Protein release from 2% (w/w) loaded microfibers

Microfiber Molecular Actual Encapsulation First phase Second (lag) Third Total weight loading efficiency duration phase phase released

(kDa) (%) (%) and release duration and duration (%) release and release

PLGA

INS 5.8 0.83 41.6 14 days 14-42 days 42-63 days 81.3

57.9% 8.7% 14.7%

LZ 14.3 0.85 42.7 21 days 21-63 days None 47.5

40.4% 7.1%

BSA 66.0 0.61 30.5 21 days 21-63 days None 40.5

28.1% 12.3%

PLLA

INS 5.8 1.46 73.0 14 days 14-35 days 35-63 days 18.5

13.4% 0.7% 4.4%

LZ 14.3 1.39 69.2 14 days 14-63 days None 8.4

4.5% 3.9%

BSA 66.0 1.24 61.8 14 days 14-63 days None 5.1

3.4% 1.7%

0.9684 to 0.9865. Cumulative protein release (%) of encapsulated proteins from PLGA and PLLA fibers at days one and day 38 after fabrication as a function of protein molecular weight is shown in FIG. 15 panels C-F. Release kinetics from PGLA fibers was observed to decrease exponentially with increased protein molecular weight {R² = 0.91-0.98). PLLA fibers having encapsulated proteins exhibited sustained release rates independent of molecular weight during the course of 38 days.

Table 6. Release parameters of protein-loaded microfibers

Microfiber Korsmeyer-Peppas

k(f") n R²

PLGA

INS 0.1281 0.2805 0.9865

LZ 0.0461 0.3430 0.9730

BSA 0.0325 0.3419 0.9868

PLLA

INS 0.0252 0.2708 0.9862

LZ 0.0137 0.3180 0.9653

BSA 0.0131 0.2441 0.9684

Example 23. Fourier transform-infrared (FT-1R) spectroscopy analysis

FT-IR analysis was performed after microfiber fabrication and after 63 days of incubation using a Perkin Elmer (Wellesley, MA) Spectrum Once B spectrophotometer with a zinc-selenide (ZnSE) universal attenuated total reflectance (ATR) attachment. Microfibers (2-3 mg) were analyzed in the region between 4000-650 cm^"1 with a resolution of 2 cm^"1. FT-IR spectra were taken at the beginning and end points of release to characterize the relative degradation of incubated microfibers. Average spectra from PLGA and PLLA microfiber formulations after fabrication (0 days) and at 63 days of incubation are shown in FIG. 16. The FT-IR spectra remained unchanged throughout the degradation analysis. The three observed absorption peaks characteristic of the Ester bonds are the ester aliphatic C=0 stretch at 1750 cm^"1, aliphatic ester C-O stretch at 1260 cm^"1, and the v(C-C) mode of the C-COO of the polymer chain at 860 cm^"1. These peaks were observed to be undiminished at 63 days of incubation. No appearance of peaks indicative of lactic and glycolic acid i.e. the carboxylic acid v(C=0) stretch at 1700 cm^"1 and the wide v(O-H) bend at 3200 cm^"1 was observed;

Example 24. Microencapsulation efficiencies

Microencapsulation efficiencies of protein loaded microfibers were determined (Table 5). PLGA formulations were observed to have lower average loading efficiencies of 38.3 ± 6.7% compared to PLLA efficiencies of 68.0 ± 5.7%. Protein loss is attributed to protein adsorption to conical tubes during the micronization process and differences in spin dope precipitation strength during phase inversion. PLLA solutions precipitated visibly faster than PLGA solutions, accounting for increased encapsulation efficiency among PLLA formulations. A reduced protein loss is envisioned in scale-up to larger batch sizes.

Example 25. Single step fabrication of wet spun binary phase composite microfibers

A single step method to fabricate a polymeric microfiber delivery system for controlled delivery of therapeutics and minimal initial burst is described using phase separated binary blends of PLLA and PLGA (75:25 ester terminated), which were wet spun by phase inversion. Polarized light microscopy, DSC, scanning electron microscopy and fluorescence microscopy were used to assess formation of composite microfibers.

To determine the conditions for phase separation of PLLA and PLGA solutions, 20% w/v 1 :1, 1 :2, and 1 :3 w/w ratios of each of the polymers were made in DCM. Solutions were mixed for 1 h and were incubated to equilibrate for 24 h. The appearance of two phases was observed by a difference in color between the polymer solutions. Films were cast from the liquid-liquid phase separated solutions. Upon verification of liquid phase separation of polymer solutions by polarized light microscopy PLLA/PLGA microfibers were wet spun by phase inversion. Polymers were dissolved in DCM in appropriate w/w ratios. The phase separated polymer solution was loaded into a glass syringe fitted with a 22 gauge spinneret and placed in a syringe pump. Since DCM is miscible with petroleum ether, the immersion of the spinneret into the coagulation bath resulted in the continuous precipitation of monofilament microfibers.

A typical micrograph of a phase separated PLLA:PLGA solution is shown in FIG. 17, in which the micrograph PLLA spherulites are visible within the continuous PLGA phase. The phase separation of 1 : 1 , 1 :2 and 1 :3 PLLA:PLGA solutions was observed under cross polarized optical microscopy. The 1 : 1 spin dope solution was thus determined to be suitable suitable for wet spinning. The inherent viscosity and molecular weight of 1 :2 and 1 :3 phase separated solutions was found to be too low to fabricate continuous monofilaments within the spin bath.

DSC thermograms exhibited the presence of two glass transitions in the first heating scan of PLLA:PLGA (1 : 1) microfibers (FIG. 18 panel A). The presence of two glass transitions, each representative of PLLA and PLGA controls, is evidence of a phase separated blend. If the solution were a miscible blend of PLLA and PLGA, the thermal properties would be between those of the two unblended polymers. Analysis of cumulative release kinetics of PLGA, PLLA and PLLA: PLGA (1 : 1 ) microfibers loaded with BSA showed that fibers with PLLA:PLGA (1 : 1) exhibited a reduced burst effect compared to fibers of PLGA or PLLA (FIG. 21 panel A).

No apparent differences in the cross-sectional morphology of binary phase PLLA:PLGA (1 : 1 ) microfibers were observed compared to fibers of PLLA and PLGA (FIG. 19). FITC- dextran was encapsulated to analyze the ability to localize potential therapeutics within binary phase composite microfibers. Localization of FITC dextran within the PLLA phase of the composite microfiber was observed by fluorescence microscopy (FIG. 20).

The ability to localize drug to an inner core of composite microfibers is useful for controlling spatiotemporal release of therapeutics from small diameter monofilament delivery systems. Composite microfibers, unlike double-walled microspheres, do not exhibit distinct core-shell morphology. Double-walled microspheres are made by solvent removal and solvent evaporation, a process that takes several hours before microspheres are fully precipitated. Wet spinning by phase inversion is a nearly instantaneous process controlled by solvent and nonsolvent miscibility. Phase separated spin dope solutions did not fully separate during microfiber precipitation and PLLA regions were solidified with the continuous PLGA phase (FIG. 21 panel B, drawing to the right).

Example 25. Fiber spinning

Biodegradable fibers were spun by phase inversion using a wet spinning system (FIG.

22). In wet spinning, an initially homogeneous polymer solution is extruded into a coagulation bath that induces phase inversion by counter-diffusion of solvent and nonsolvent (FIG. 22 panel A). As a result phase separation is initiated of the polymer solution into two phases, a polymer lean and polymer rich phase. Phase separation of polymers continues during the period of time the fiber is incubated in the coagulation bath. This time duration is referred to as the residence time. In wet spinning the parameters of size, shape, morphology, and strength of wet spun filaments, contribute to mechanical strength and drug delivery attributes, and these parameters depend on factors including polymer concentration, solubility of polymer in solvent, solvent/nonsolvent miscibility parameters, residence time, and fiber drawing methods.

To spin fibers, spin dope solutions of polymer formulations dissolved in DCM were added to a 5 mL pump-controlled syringe fitted with a 22-gauge spinneret. Spin dope solutions were extruded into a nonsolvent coagulation bath resulting in the rapid de-solvation of liquid polymer streams and continuous formation of monofilaments (FIG. 22 panel A). Different spin dope solutions were investigated. For encapsulation of DXM, 10% and 20% (w/v) solution compositions were fabricated from blends of PLLA, PLGA and PVP in co-solvent mixtures of 6: 1 (v/v) DCM to THF. A co-solvent mixture was used to increase solubility of DXM in spin dope solutions. Fibers were extruded at 0.02-1.2 mL min^"1 into petroleum ether, 2-propanol, or mixtures thereof and were collected from the spin bath using one of the following methods: fibers were either left in the bath throughout the extrusion process (as-spun); drawn from the bath during extrusion (solution-drawn); or were removed from the spin bath and wound around metal bobbins under tension (post-drawn). Each of these processes was developed to determine the applicability of wet spun fibers for designing hybrid devices.

Example 26. Multifilament yarn production

Single fibers were used to produce multifilament yarns. Monofilaments were fabricated by extruding spin dope solutions of PLLA, PLGA, and PVP into a 50:50 (v/v) mixture of 2- propanol to petroleum ether coagulation bath for 1.5 min. Addition of 2-propanol was determined to be suitable for some formulations to overcome coagulation around the spinneret tip by slowing the rate of precipitation for the formation of continuous fibers. Fibers were cured for four minutes, and were removed from the spin bath and placed in an empty beaker. Samples were untangled and wound around metal bobbins (post-drawn) on a rotating mandrel positioned above the beaker. To create 6-ply yarns, six bobbins, four monofilaments of a 10% (w/v) polymer composite formulation 'A' and 2 monofilaments of a 20% (w/v) composite formulation 'B' were placed on a winding mandrel (pattern B, A, A, A, A, B). Monofilament fibers were grouped together using a surgical clamp and twisted along the longitudinal axis in the 'Z' direction (filament inclination from top right to bottom left) to create 30-inch multifilament yarns (FIG. 22 panel B). Yarns were wound around an empty bobbin and stored at -20°C. Example 27. Scanning electron microscopy of monofilaments of different wet spun formulations

Representative samples of each formulation were evaluated by scanning electron microscopy. Samples were mounted on adhesive metal stubs and sputter-coated with a 50-100 A layer of gold-palladium (Emitech, Kent, England). Micrographs were taken with a Hitachi S- 2700 (Tokyo, Japan) at an accelerating voltage of 8 kV using a Quartz PCI digital imaging system. To determine the average diameter of microfilaments and yarns, five fields of view were obtained at lOOx and 35x, respectively. The average porosity of cross-sectioned fibers was also determined from five fields of view at l ,300x. Micrographs were analyzed using NIH ImageJ software (Bethasda, MD). Porosity was calculated as the pore area divided by the total cross- sectional area. Results of scanning electron microscopy of wet spun monofilaments of different formulations showed diverse surface features (FIG. 23).

Table 7. Wet spun fiber formulations with DXM loadinj

Polymer composition (% w/w)

PLLA PLGA PLLAo PVP

Solution-drawn 10% solution

SF- 10a 95 5 - -

SF- 10b 94 5 - 1

SF- 10c 95 - 5 -

SF- lOd 94 - 5 1

SF- lOe 99 - - 1

As-spun 20% solution

SF- 20a 50 50 - -

SF- 20b 67 jj - -

SF- 20c 75 25 - -

SF- 20d 80 20 - -

Post-drawn 10% solution

SF- 10 formulation 'A' 94 5 - 1

Post-drawn 20% solution Example 28. Dexamethasone release analysis

The polymer compositions of DXM-loaded formulations are shown in Table 7.

Monofilaments and multifilament yarns (triplicate samples, 10 mg each) were incubated in 1.4 mL of phosphate buffered saline (PBS) at 37°C. At each time point, 1.0 mL of the releasate was removed and replaced with fresh PBS. Amounts of DXM released from individual fibers and multifilament yarns were quantified by UV absorption at 239 nm. A 96-well quartz plate was used to minimize background at the reading frequency. Drug concentration was determined by comparing absorbance values to standards of known DXM concentration. Polymer composition was observed to modulate drug release kinetics of DXM- loaded monofilaments (FIG. 26).

Example 29. Engineering fibers with desired release kinetics from multifilament yarns

The release kinetics of multifilament yarns were predicted from the experimental release of individual monofilaments. For a multifilament yarn, the theoretical release as a function of release time, t was calculated according to

Theoretical Release (t)

+/BRB(V (7) in which / is the fraction of fibers from any given formulation relative to the total number of discrete monofilaments included in the yarn, and R is the amount of drug released from individual formulations as a function of release time, t. This equation is applied to number of individual formulations twisted to create multifilament drug-eluting yarns.

Example 30. In vitro bioactivity of dexamethasone released from multifilament yarns

Human aortic valve interstitial cells (hVICs) isolated from human cryopreserved conduits passage 2 through 4, (from the Cardiac Regenerative Surgery Research Laboratory, Children's Mercy Hospital) were used for in vitro analyses. Cells were maintained in hVIC media, Dulbecco's Modified Eagle's Medium (DMEM F12, Gibco) supplemented with 10% fetal bovine serum (FBS, Gibco), 1% L-glutamine, 1% penicillin/streptomyocin, and 1% amphotericin B. The biological activity of eluted DXM was determined using a cell proliferation assay (Reil TD et al, J Surg Res, 1999;85: 109-114). Passage 5 hVICs were plated with 10% FBS growth media at 5000 cells/cm² into collagen type 1 coated 12-well plates (BD BioCoat, San Jose CA) and cultured for 24 h. Cell growth was arrested by washing plates with PBS and adding 1 % FBS serum starvation media. Growth arrest was maintained for 48 h to allow for cell cycle synchronization. Synchronized cells were then re-stimulated with 10% FBS growth media containing 10^"7 mol/L DXM. DXM-supplemented growth media was made using eluted DXM releasates from multifilament yarns at 1-day and 56-day incubation. Fresh, unencapsulated drug and PBS vehicle were used as controls. The media for each experimental group was changed every 24 h continuing to a total exposure of 72 h. Results of the bioactivity of DXM released from multifilament yarns are shown in FIG. 28.

Example 31 Mechanical testing Uniaxial tensile tests were conducted using a materials testing system (Instron Model 4442) in accordance to the United States Pharmacopeia (USP) absorbable suture testing standards (United States Pharmacopeia and National Formulary (USP 34-NF 29), United States Pharmacopeial Convention, Rockville, MD, 2010). Monofilaments and yarns were secured to a paper frame (25 mm x 25 mm) and loaded into the crosshead clamps of the machine. Prior to loading, the sides of the paper template were cut and the sample was maintained intact. An elongation rate of 50 mm min^"1 was applied until failure, and load-displacement data was collected by digital acquisition system. Load at break is a measure of elongation until fiber fracture. The ductility or strain at failure (ε = ALLo-l) is calculated from the original gauge length (Lo) and the change in gauge length (AL) as recorded by crosshead movement until fiber fracture. Ten replicates of each formulation were analyzed. Results of mechanical testing are shown in Table 8.

Example 32. Statistical analysis

Statistical analysis was performed using SPSS v.19 statistical software (Chicago, IL). All data are expressed as mean ± S.D. An ANOVA and a post hoc Tukey multiple comparisons test were used for group comparisons. To determine the significance of theoretical multifilament yarn drug release in comparison to experimental release, a two-tailed t-test was performed. A p value of less than 0.05 was considered statistically significant.

Example 33. Morphology of wet spun monofilaments

Fibers were fabricated by applying post-handling techniques, varying spin dope precipitation strengths, and altering the solvent/nonsolvent miscibility parameters and were observed to have diverse surface structures. For example, as-spun PLGA fibers extruded into petroleum ether were characterized by a smooth surface morphology (FIG. 23 panels A and B). Conversely, solution-drawn PLLA₀ 4 fibers extruded into a mixture of 2-propanol and petroleum ether and stretched under tension as the dope solution precipitated were characterized by longitudinal grooves along the direction of fiber drawing (FIG. 23 panels C and D). Fibers spun from high and low solution concentrations and extruded under otherwise identical conditions also resulted in diverse surface architectur e. Initiation of gelation of fibers wet spun from 10%

(w/v) PLLA (FIG. 23 panel E) was observed to be slower than that of 15% (w/v) PLLA (FIG. 23 panel F) fibers since there was more solvent volume to diffuse into the petroleum ether coagulation bath. Similarly, changing the molecular weight of the polymer solution was observed to alter the surface morphology of wet spun filaments. High molecular weight PLLA (FIG. 23 panel G) solutions were characterized by smoother surface structures compared to lower molecular weight PLLA₀.94 (FIG. 23 panel H) solutions spun under identical conditions. Altering the nonsolvent bath composition was observed to result in a change in the surface roughness of wet spun filaments. Composite spin dope solutions extruded into petroleum ether displayed smooth surface topography (FIG. 23 panel I), and the same solution extruded into a 50:50 mixture of 2-propanol to petroleum ether possessed surfaces with micron-rough grooves and spherical protrusions (FIG. 23 panel J). Solutions precipitated visibly slower in 2-propanol than petroleum ether. The changes observed in the rate of precipitation were a direct result of the solvent system used. The solubility parameter of DCM (5DCM = 20.2 MPal /2) is 5.5 units higher than petroleum ether (PE) (δΡΕ = 14.8 MPal/2) and 3.6 units from 2-propanol (2-P) (52- P = 23.8 MPal/2), indicating that petroleum ether is a better nonsolvent for the polymer. Thus, the onset of precipitation was faster with petroleum ether and a smooth surface morphology was obtained.

Composite formulations selected for multifilament yam production further demonstrated the diverse effects of processing conditions on surface structure and cross-sectional morphology of wet spun filaments. Post-drawn 10% (w/v) composite fiber surfaces were observed to have longitudinal striations with many spherulites (FIG. 24 panels A and C) compared to 20% (w/v) composite fibers having nano-porous surfaces and fewer spherulites (FIG. 24 panels E, G). The cross-sectional porosity of composite fibers was also significantly different. The porosity of 10% and 20% (w/v) composite fibers was observed to be 3.0 ± 1.8 μιη² and 9.1 ± 2.8 μηι², respectively.

Example 34. Drug release kinetics of as-spun and solution-drawn monofilaments

To investigate the effect of polymer composition and concentration on DXM release, the kinetics of drug release from composite monofilaments of 10% and 20% (w/v) polymer concentrations were determined. Polymer concentrations were selected based on mechanical characteristics of blank as-spun PLLA fibers (FIG. 25). Initial release analyses were terminated at 35 days in vitro to mimic the timeline of acute and chronic phases of wound healing.

The cumulative DXM release from 10% (w/v) composite monofilaments loaded with 2.5% (w/w) DXM is shown in FIG. 26 panel A. Each formulation exhibited an initial burst release from the first day of incubation, followed by a continuous linear release to 35 days. The initial release of drug was fastest from fibers prepared with more hydrophilic polymers.

PLLA PLGA/PVP composite fibers (SF-lOb) were observed to have the highest initial burst of 5.5 ± 0.7 μg mg^"1 of fibers, and the other formulations each exhibited a similar initial drug burst release of 2.5 ± 0.5 μg mg^"1 of fibers. The release kinetics of 20% (w/v) composite fibers loaded with 1.5% (w/w) actual DXM are shown in FIG. 26 panel B. Composite fibers were frmulated from ratios of 1 : 1 (SF-20a), 1 :2 (SF-20b), 1 :3 (SF-20c), and 1 :4 (SF-20d) PLGA to PLLA (20-50% PLGA content). Fibers prepared from 20% (w/v) spin dope solutions were observed to exhibit significantly different release kinetics compared to fibers wet spun from 10% (w/v) polymer solutions. PLGA/PLLA formulations prepared from 20% (w/v) spin dope solutions exhibited very little burst (<3% encapsulated DXM) within the first day of incubation, with a presence of PLGA to the extent of 50% fiber composition. Each formulation was observed to have a similar lag phase to 14-day incubation. At this phase, the formulation having 50% PLGA content showed a marked increase in release rate. At 35 days, the cumulative drug release was 13.6 ± 2.7 μg mg^"1 of fibers (78% encapsulated DXM) for 1 : 1 ; 2.3 ± 0.2 μg mg ¹ of fibers (21% encapsulated DXM) for 1 :2; 1 .7 ± 0.7 g mg ¹ of fibers (12% encapsulated DXM) for 1 :3; and 1.8 ± 0.8 μg mg^"1 of fibers (12% encapsulated DXM) for 1 :4 PLGA/PLLA formulations. Drug release from 20% (w/v) polymer solutions with 50% PLGA content was observed to significantly increase compared to amount of the initial burst. PLLA/PLGA/PVP (SF- 10b) and PLLA/PLGA (SF- 10a) exhibited the greatest overall drug release (50% and 60% encapsulated DXM) with average daily release rates of 1.1 μg mg ¹ and 1.8 μg mg^"1 of fibers, respectively. PLLA/PVP (SF-lOe), PLLA/PLLA₀ VP VP (SF- lOd), and PLLA/PLLA0.94 (SF-lOc) formulations released 24-29% total encapsulated drug and were observed to have a similar daily release rates of 0.6-0.7 μg mg^"1 of fibers.

Example 35. Drug release kinetics of post-drawn composite monofilaments and 'Z' twisted multifilament yarns

Drug-eluting wet spun microfilaments were engineered into macro-structured implants. Crystalline regions of PLLA were observed herein to contribute to mechanical strength, and to potentially decrease the rate of drug release. A linear release of up to 28% total encapsulated

DXM was observed (Example 12, FIG. 27) from PLLA fibers incubated for eight weeks in vitro, and therapeutic levels of DXM from wet spun fibers using a low (7.5%) PLLA concentration were achieved. To increase the rate of release of DXM from higher PLLA concentrations (10% and 20% w/v) composite fibers with the addition of a lower molecular weight PLLA, amorphous PLGA, and water-soluble polymer PVP were formulated.

Monofilaments used for yarn fabrication were spun from 10% and 20% (w/v) concentrations. Polymer blends were selected to achieve therapeutic ranges of DXM treatment (Table 7). The PLGA and PVP content was increased for formulation of the 20% (w/v) composite fibers to further modulate release. The release kinetics of post-drawn monofilaments are shown in FIG. 27 panel A. Monofilaments spun from 10% (w/v) composite fibers loaded with 2.8% actual DXM were observed to have substantially less burst release compared to 20% (w/v) polymer solutions formulated with 2.4% actual DXM.

After the initial burst phase, fibers made from 10% (w/v) fibers exhibited linear release rates (R² = 0.98) in comparison to 20% (w/v) solutions, which exhibited logarithmic drug release (R² = 0.96) up to 56 days in vitro. The average daily release after the initial burst for composite solutions was 0.7 μg mg^"1 for 10% (w/v) fibers. For 20% (w/v) fibers, the average daily release up to 21 -day incubation was 1.8 μg mg^'1, followed by a release rate of 0.2 μg mg^"1. At 56-day incubation, 10% (w/v) formulations released 1 1.6 ± 0.4 g mg^"1 of fibers (21% encapsulated DXM), and 20% (w/v) formulations released 17.2 ± 0.4 μg mg^"1 of fibers (36% encapsulated DXM).

Post-drawn fibers are potentially useful for tailoring release of therapeutics from medical implants by the formation of multifilament yarns. Yarns produced from several different combinations of single fibers are envisioned to achieve desired release profiles for specific clinical applications including sequential release of multiple small molecules or therapeutics. FIG. 27 panel B shows release profiles predicted from the experimental release of single fibers in FIG. 27 panel A using equation 7 for five different theoretical 6-ply yarns. The experimental DXM release from a 6-ply multifilament yarn formulation was evaluated. The release kinetics of 6-ply multifilament yarns produced by 'Z' twisting four single filaments of formulation 'Α', with two single filaments of formulation 'B' (Table 7) were observed to have release profiles within the ranges of individual filaments (FIG. 27 panel C). Yarns released 5.3 μg mg^"1 (15% encapsulated DXM) in an initial burst release after one day of incubation, followed by an average daily release of 0.8 μg mg^"1 from day 1 to day 56 incubation. By day 56 of incubation, yarns released a total of 12.4 ± 1 mg^"1 (35% encapsulated DXM). Good agreement was observed between the experimental drug release and the theoretical prediction. Release samples at the measured time points except days 14 and 28 were not statistically different (p < 0.05) than the predicted release, suggesting monofilament twisting did not affect the release kinetics of constituent monofilaments. Therefore, with respect to drug release, fabrication of multifilament yarn from monofilament yarn did not affect linearity of release. Example 36. Biological activity of eluted dexamethasone from multifilament yarns

The effect of dexamethasone released from multifilament yarns on the proliferation of hVICs was studied to evaluate the biological activity of eluted drug. The proliferation of hVICs cultured in the presence of DXM-supplemented media from multifilament yarns that had been incubated for 1 day, and 56 days in comparison to control media was determined (FIG. 28). The hVICs cultured in the presence of medium that contained DXM proliferated at a significantly reduced rate after 72 hours compared to control cells with PBS vehicle. Proliferation of hVICs grown in medium that contained either fresh or eluted DXM were similar. These findings confirmed that wet spinning did not alter the biological activity of encapsulated DXM since each of eluted and unencapsulated DXM treatment was observed to affect the growth and proliferation of hVICs to the same extent.

Example 37. Mechanical properties of post-drawn monofilaments and 'Z' twisted multifilament yarns

The tensile properties of wet spun fibers wound around metal bobbins were evaluated to analyze the effects of post-drawing and multifilament yarn production on filament breaking strength and ductility. A summary of the tensile properties of single filaments (SF) and twisted (TW) 6-ply yarns is shown in Table 8. After fabrication, blank and drug-loaded monofilaments were observed to have a slight increase in load to failure as a function of increasing the polymer concentration from 10% to 20% (w/v). Addition of DXM into wet spun monofilaments did not significantly reduce the breaking strength of fibers, unlike other studies encapsulating drugs within wet spun fibers (Chang HI et al., J Biomed Mater Res Part A, 2008;84:230-237; Mack BC et al., J Control Release, 2009; 139:205-21 1 ; Rissanen M., J Appl Polym Sci,

2010;1 16:2174-2180). The ductility of 10% and 20% (w/v) blank (p = 0.061) and drug-loaded (p = 0.155) were similar. DXM-loaded and blank multifilament yarns possessed a fivefold increase (p = 0.000) in load bearing capacity compared to individual DXM-loaded and blank fibers. The marked increase in load at failure correlated well with the summation of the failure strength of individual monofilaments. The ductility of yarns was significantly reduced (p <0.05) compared to individual blank and drug-loaded fiber formulations, indicating that monofilament twisting removed a portion of the strain within constituent monofilaments.

Table 8. Summary of tensile properties of post-drawn monofilaments and 'Z' twisted multifilament yarns. Data are represented as mean ± S.D. («=TQ)

Sample type Diameter (μιη) Load at failure (mN) Strain at failure (mm/mm)

SF- lObknk 49.5 ± 7.9 64.3 ± 7.7 2.10 ± 0.59

SF- 20blank 70.2 ± 7.9 89.9 ± 8.7 1.48 ± 0.72

TWbknk 152.1 ± 6.3 341.5 ± 27.1 1 .34 ± 0.35

SF- 10 48.4 ± 6.6 71.0 ± 8.7 2.55 ± 0.51

SF- 20 64.6 ± 6.9 77.3 ± 3.6 2.12 ± 0.59

TW 153.9 ± 7.5 447.1 ± 40.9 1.16 ± 0.40 Example 38. Handling capabilities of 'Ζ' twisted multifilament yarns

To scale up of microfilaments into complex structures, single microfilaments were grouped together and twisted into multifilament yarns. Six fibers were twisted in the 'Z' direction for the formation of 6-ply multifilament yarns. Twisted fibers were scaled to a thickness of around 150 μηι with even entwining and small spaces between single fibers (FIG. 22 panel B). The handling capabilities of multifilament yarns was observed to have improved compared to constituent monofilaments. Yarns were capable of being readily braided, knitted or woven into complex geometries (FIG. 29)

Claims

What is claimed is:

1 . A wet spun microfiber composition comprising at least one polymer wherein the composition comprises:

a porous polymeric microstructure, and further comprises at least one encapsulated therapeutic agent, wherein the therapeutic agent is located in the microstructure and is controllably releasable from the composition, and wherein the microfiber has a degree of crystallinity at least 10% greater than that of control polymer prior to wet spinning.

2. The composition according to claim 1 having a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

3. The composition according to claims 1 or 2, wherein an encapsulated first therapeutic agent comprises dexamethasone. 4. The composition according to claims 1-3, further comprising an encapsulated second therapeutic agent.

5. The composition according to claim 4, wherein the second therapeutic agent comprises at least one selected from the group: a drug; a protein, for example, Nog (Noggin); a peptide; a sugar; a carbohydrate; and a nucleotide sequence.

6. The composition according to claim 5, wherein the protein comprises at least one selected-from the group: a growth factor, an immunoglobulin, an enzyme, and a peptide antibiotic.

7. The composition according to claim 5, wherein the nucleotide sequence comprises a vector.

8. The composition according to claims 1-7, wherein the polymers are at least one of poly- 1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP).

9. The microfiber composition according to claims 1 -4, wherein the composition comprises at least about 75% of the initial tensile strength for at least about five weeks.

10. A method of producing a wet spun microfiber composition having a porous multi-layer polymeric microstructure comprising the steps of:

mixing at least one polymer and at least one therapeutic agent with a solvent to form a solution; and

wet spinning the material by phase inversion, thereby producing the microstructure, wherein the composition has a degree of crystallinity which is at least 10% greater than that of control polymer prior to wet spinning.

1 1. The method according to claim 10, wherein a first therapeutic agent is dexamethasone.

12. The method according to claim 10, further comprising mixing the polymer solution with a second therapeutic agent prior to wet spinning.

13. The method according to claim 12, wherein the second therapeutic agent is selected from the group of: a protein a peptide, a sugar, a carbohydrate, a nucleotide sequence, and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anti-coagulant; an antitumor; an anti-viral; an anti-bacterial; an anti -mycobacterial; an anti-fungal; an antiproliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a non- steroidal anti-inflammatory agent (NSAID).

14. The method according to any of claims 10-13, wherein the solvent comprises at least one of dichloromethane and tetrahydrofuran. 1 . The method according to any of claims 10-14, wherein wet spinning comprises loading the material into a syringe, and dispensing the material into a coagulation bath, wherein the coagulation bath comprises a non-solvent, thereby obtaining phase inversion.

16. The method according to claim 15, wherein the method further comprises selecting the solvent and the non-solvent having different solubility parameters, wherein the difference between the solubility parameters affects the rate of solidification of the polymer, the extent of solvent induced crystallization of the polymer, and the degree of crystallinity of the composition.

17. The method according to claim 15, wherein the coagulation bath comprises petroleum ether.

18. A method of treating a subject having a medical condition comprising, contacting the subject with a wet spun microfiber composition comprising at least one polymer wherein the composition comprises a porous polymeric microstructure, and further comprises at least one encapsulated therapeutic agent, wherein the therapeutic agent is located in the microstructure and is controllably releasable from the composition, and wherein the microfiber has a degree of crystalhnity at least 10% greater than that of control polymer prior to wet spinning. 19. The method according to claim 18, wherein a first therapeutic agent is dexamethasone.

20. The method according to claim 18, further comprising an encapsulated second therapeutic agent. 21. The method according to any of claims 18-20, wherein the composition has a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.

22. The method according to any of claims 18-21, wherein the polymers are at least one of poly- 1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP).

23. The method according to any of claims 1 8-22, wherein the medical condition is at least one selected from the group of: a burn, an abrasion, a laceration, a pathology, a cancer, and an infection. 24. The method according to any of claims 20-23, wherein the second therapeutic agent comprises at least one selected from the group: a sugar; a carbohydrate; a nucleotide sequence; a protein selected from the group of: a growth factor, an immunoglobulin, an enzyme, and an antibiotic; and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anti-coagulant; an anti-tumor; an anti-viral; an anti-bacterial; an anti-mycobacterial; an anti- fungal; an anti-proliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a non-steroidal anti-inflammatory agent (NSAID).

25. The method according to claim 24, wherein the nucleotide sequence comprises a vector.

26. A kit for treating a subject having a medical condition comprising:

a wet spun microfiber composition having at least one polymer wherein the composition comprises a porous polymeric microstructure, and further comprises at least one encapsulated therapeutic agent, wherein the therapeutic agent is located in the microstructure and is controllably releasable from the composition, and wherein the microfiber has a degree of crystallinity at least 10% greater than that of control polymer prior to wet spinning;

instructions for use; and,

a container. 27. The kit according to claim 26 wherein a first therapeutic agent is dexamethasone.

28. The kit according to claim 26 further comprising a second therapeutic agent.

29. The kit according to any of claims 26-28, wherein the polymers are at least one of poly- 1 -lactic acid (PLLA), poly-lactic-co-glycolide (PLGA) and polyvinylpyrrolidone (PVP).

30. The kit according to any of claims 28-29, wherein the second therapeutic agent comprises at least one selected from the group: a sugar; a carbohydrate; a nucleotide sequence; a protein selected from the group of: a growth factor, an immunoglobulin, an enzyme, and an antibiotic; and a drug, for example, an anti-apoptotic; an immunosuppressant; a pro-apoptotic; an anti-coagulant; an anti-tumor; an anti-viral; an anti-bacterial; an anti-mycobacterial; an antifungal; an anti-proliferative; and an anti-inflammatory, for example, a steroid selected from the group of: a cortisone compound, for example a dexamethasone, and a sex-related hormone; and a non-steroidal anti-inflammatory agent (NSAID).

31. The kit according to claim 30, wherein the nucleotide sequences comprises a vector.

32. The kit according to any of claims 26-3 1 wherein the composition has a structure selected from the group of: a fiber, a suture, a sphere, an implant, and a scaffold.