Methods and compositions to administer therapeutics to soft tissue surfaces

Abstract

A dry drug delivery system comprising a middle layer encased in two outer layers, the middle layer and the two outer layers each comprising a mixture of electrospun fibers. Also disclosed is a water-dissolvable dry device containing a physical mixture of a first fiber and a second fiber in which the first fiber and the second fiber are configured to react with each other and with a biological tissue upon hydration, and the device transforms into a hydrogel upon contact with the biological tissue. Further provided is a method for delivering a bioactive substance to a soft tissue using the system or device.

Description

METHODS AND COMPOSITIONS TO ADMINISTER THERAPEUTICS TO SOFT TISSUE SURFACESRELATED APPLICATION

[0001] This application claims the benefit of US Provisional Application Serial No. 63 / 551,400 filed on February 8, 2024. The entire teachings of the above application is incorporated herein by reference.BACKGROUND

[0002] Defining the Challenges

[0003] Issue of Low Encapsulation and Degradation of Labile Molecules, Irritation, Low Retention on Tissue and Rapid Release Kinetics of Compounds

[0004] Low Encapsulation due to Degradation of Labile drug: Highly water- soluble small molecule (synthetic and biologically derived) compounds and larger molecules such as peptides, proteins, and nucleic acids are difficult to formulate in matrices that can encapsulate high concentrations of the labile drug (like a protein or DNA) and achieve a sustained release profile of intact drug.

[0005] These molecules (whether small or large) can be labile, prone to degradation in certain microenvironments (pH, salts, enzymes, organic / aqueous interface, solvents), environmental conditions (high temperatures, etc.), physical shear forces, ionic interactions and exposure to extreme conditions. Strategies to formulate these approaches have involved selection of optimum pH and buffer counterions, identification of excipients, and blends of excipients that assist in the stabilization of these labile molecules in solution form, and in solid form. When formulated withcertain excipients, labile molecules can be protected from proteases and micro- environmental stresses, but high concentrations are extremely difficult to achieve.

[0006] In addition, for topical tissue surfaces, sustained release formats have issues of retention on the target site. For topical tissue surfaces (e.g., skin, sublingual, intranasal, vaginal, and ocular), a sustained release "film-type" format is a preferred strategy. For tissue spaces like ocular mucosa, a film dosage format is an insert. In cavity-type tissues such as intranasal, sublingual, rectal and buccal tissues, delivery of a medication that requires a sustained presence is challenging, due to the rapid turnover of fluid. Another tissue type is the esophageal mucosa, which requires retention on-site for site-specific sustained administration of medication. Vaginal, rectal, intranasal and urological mucosal tissues can benefit from sustained release from inserts, particularly where the force of gravity can extrude any viscous solutions, gels, and ointments. A well-adherent insert that will adhere to the tissue (mucosal and non-mucosal) and resist extrusion by gravity can achieve site-specific targeted dosing, without waste of medication.

[0007] It is challenging to incorporate highly water-soluble compounds into polymeric film inserts in sustained release formats, with the compound dispersed in polymeric matrices. For inserts that are less than 1 mm in thickness, the diffusive path length for an encapsulated compound is small, resulting in rapid drug release. Rapid release kinetics is observed for both hydrophobic and hydrophilic molecules, due to small diffusive path lengths.

[0008] Further, another challenge arises when labile active pharmaceutical ingredients (API), or mixtures of pharmaceutical ingredients are required to be encapsulated in their intact forms, within the polymeric delivery system, and should stay intact, or maintain a high degree of integrity during storage and during release from the drug delivery system matrix. The components of the delivery system cannot react with the active ingredient, potentially compromising its biological potency. This class of labile molecules normally fall into the category of biomolecules such as antibodies, proteins, peptides, aptamer and polynucleotides such as DNA. Additionally, compounds that are susceptible to degradation by hydrolysis, or oxidation or another mechanism can benefit from being encapsulated in a matrix composition that not only sustains its release but keeps it stable from elements ofdegradation such as proteases, nucleases, oxidative, hydrolytic and other degradative species. Thus, all components of the delivery system must be chemically compatible with the encapsulated compound and play a protective role toward maintaining integrity.

[0009] Lastly, film inserts that are placed on the target tissues should be removeable once the drug is depleted from the matrix. Inserts that are placed on the tissue should also be dissolvable-dissolving after drug release, or in the case of "nodrug" inserts, post-placement. The insert should be tissue-biocompatible and dissolution of the components should be modulable.

[0010] Issue of Irritation

[0011] The fundamental challenge with both implantable and topical drug delivery systems is inherent: irritation caused by a non-favorable molecular interaction of the insert with underlying tissue, particularly if the tissue has water content, is mucosal and has nerve endings. Irritation caused by inserts at the local tissue is a common foreign body response. One of the most common issues is with inserts that are placed on mucosal tissue, causing an unpleasant sensation of irritation and causing redness, local discomfort, pain and erythema. However, certain biomaterials (PLGA, PLA, PCL, and copolymers and blends) comprising the inserts become stiff, contract into devices with sharp edges, lose malleability and flexibility under physiological temperature and pH. Stiff, inflexible inserts become highly irritating to underlying tissue, especially in the eye, when a previously flexible insert feels like shards of glass within hours. PLGA degrades into acidic fragments creating a dense surface buildup of acidic impurities in the local microenvironment of the implanted insert. An acidic microenvironment however transient, can result in issues with biocompatibility, inducing the generation of immune responses and manifested by mild irritation and redness, or more severe, such as burning. Due to their enhanced cellular uptake and biodistribution, PLGA particulates have been known to affect cell growth, viability and tissue responses, and clinically manifest as inflammation, sensation of irritation and redness. These issues necessitate continued development of innovative approaches to generate biocompatible tissue-insert interfaces to minimize irritation.

[0012] Issue of Retention of Inserts

[0013] Mucosa present in accessible and inaccessible spaces have the challenge of insert retention after placement. These include ocular, intranasal, rectal, vaginal and cervical, intra-uterine, intra-urological, esophageal, buccal, sublingual, pulmonary and bronchial mucosa.

[0014] One of the key performance criteria of drug delivery systems that are administered to the tissue surfaces is retention of the drug-containing device longterm. "Long-term" can be relative: any duration between a few hours to many days (e.g., 5-180 days). Electrospun inserts placed in the conjunctival fornix (lower and upper eyelids), are characterized by their high interstitial porosity to enable flux of the biological milieu thereby integrating the insert into the underlying tissue. Additionally, PLGA-based electrospun, or solvent-cast inserts lose shape, undergoing dimensional contraction. A combination of dimensional instability, and increase in stiffness of the insert, result in delamination of the device from the underlying tissue, cause local irritation due to rubbing of sharp edges against soft tissue and inevitable loss from its site of administration.SUMMARY

[0015] To overcome the challenges set forth above, a dry drug delivery system is provided that includes a middle layer encased in two outer layers, the middle layer and the two outer layers each comprising a mixture of electrospun fibers, the two outer layers each including one or more polymers selected from ethyl vinyl acetate, poly trimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) having a copolymer ratio of 90:10, PCL, and P(TMC) -polyethylene oxide (PEG)-P(TMC) in which the middle layer is formed of core-shell fibers, each of which includes (i) a core that contains a bioactive substance and one or more excipients to protect the bioactive substance from degradation, and (ii) an outer shell formed of a hydrophobic elastomeric polymer.

[0016] Also provided is a method for delivering a bioactive substance to a soft tissue of a subject by applying the dry drug delivery system to the surface of a soft tissue of a subject before or after adding a hydrating fluid to the soft tissue.

[0017] Further, disclosed is a water-dissolvable dry device that is a physical mixture of a first fiber and a second fiber in which the first fiber contains one or more of sodium hyaluronate, castor oil, polyvinyl alcohol (PVA), and polysorbate 80 or polysorbate 20, and, optionally, a buffer salt, an amino acid, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinyl pyrrolidone, dextran 70, albumin, PEG-multi-arm amine, polyethylene glycol, collagen type 1, glycerol, PEG40- stearate, and the second fiber contains PVA and sodium alginate and, optionally, sodium borate, dextran 70, PEG-multi-arm-succinimidyl ester, polypropylene glycol, propylene glycol, collagen type I, a bioactive substance, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or beta 1>6 glucan. The first fiber and the second fiber react with each other and with a biological tissue upon hydration, and the device transforms into a hydrogel upon contact with the biological tissue

[0018] The details of one or more embodiments of the invention are set forth in the description and drawings below. Other features, objects, and advantages of the invention will be apparent from the description, the drawings, and the claims.BRIEF DESCRIPTION OF THE DRAWINGS

[0019] The foregoing and other objects, features and advantages will be apparent from the following description of particular embodiments of the disclosure, as illustrated in the accompanying drawings.Fig. 1 : Exemplary Size and Shape of an Ophthalmic InsertFig. 2: Exemplary Applicator Tool DesignFig. 3: Left-top: Scanning Electron Micrograph of the Dissolvable Insert, (b) Leftbottom: Scanning Electron Micrograph of Ocular Mucosa, (c) Hydrated Insert placed on the ocular surfaceFig. 4: Young’s Modulus of Dry InsertsFig. 5: Young’s Modulus of Hydrated InsertsFig. 6: Scanning Electron Micrograph of Dissolvable, In-Situ Tissue Crosslinkable Insert (PVA- Alginate)Fig. 7: Cross-linking between borate ions and the hydroxyl groups in the polymer chains of polyvinyl alcoholFig. 8: Crosslinking Kinetics and Dissolution of Collagen-Containing Hydrogel FormulationFig. 9: Hydrogel Insert ApplicationFig. 10: Hydrogel FormationFig. 11: Application of Insert to Conjunctival FornixFig. 12: Application of Wetting SolutionFig. 13: Hydrogel Insert after WettingFig. 14: Insert Intercalated into Tissue after 1 Hour in both EyesFig. 15: Hydrogel insert applicationFig. 16: No Visible Hydrogel 24 hours after insert ApplicationFig. 17: After Insert ApplicationFig. 18: Hydrogel Formation after Rapid Uptake of WaterFig. 19: Ocular Surface Before ApplicationFig. 20: SB-NANOM-X-15-14 Insert ApplicationFig. 21: Postmortem Ocular Surface, T= 6 hoursFig. 22: Bioadhesion of Bonding occurs ~30 minutes after application of insert and hydrating buffer (Left=Peel strength of Insert from Tissue at T=0; Right=Peel Strength of Insert from Tissue at T=30 minutesFig. 23: Peptide Before and After Incorporation into InsertFig. 24: Mucosa- Adhered Annular Hydrogel- Sheath Hydrophobic Core Fiber InsertFig. 25: Small Peptide Exenatide Release from Crosslinked Dissolvable InsertsFig. 26: Small Peptide Release from Electrospun InsertFig. 27: Inserts Demonstrate High Elastomeric CharacteristicsFig. 28: IgG-Containing Drug ProductFig. 29: Release of Intact Protein from the Insert Drug ProductFig. 30: Sustained Release of IgG from Insert Drug Product with P(TMC)- LactideFig. 31: Batch Reproducibility of IgG-containing Insert ProductFig. 32: Size Exclusion Chromatograms of Protein-Containing Inserts Sterilized by Electron Beam Exposure at 25 kGy (Left) and 15 kGy (Right)Fig. 33: Dimensional Stability Comparison: PCL12 (Blend), PLGA, PDO and PCL80KFig. 34: tO (on left) and 114 days (on right) of the PCL12 Blended Composition showing the dimensional stability over time.Fig. 35: tO (on left) and t9 days (on right) showing the dimensional stability over time.Fig. 36: (A) 6 mm insert to be placed on tissue; (B) square insert to be placed on tissue, (C, D) 8 mm insertsFig. 37: Image A: Score = 0 (top left), shows no adhesion and was scored with a 0. Image B: Score = 1 (top right) demonstrates some adhesion to the tissue, peeling off, stretching the conjunctiva tissue. Image C: Score = 2 (bottom left), peels off with more resistance, the mesh starts to stretch and tare. Image D: Score = 3 (bottom right), great adhesion, sample can be lifted by mesh.Fig. 38: Scanning Electron Micrograph of a core-shell only, insertFig. 39: In-vitro release of protein (Core-sheath vs. Sandwich)Fig. 40: Sustained Release profiles of Proteins from “Sandwich” InsertsFig. 41: Scanning Electron Micrographs of the Sandwich Insert in the Preannealed, Post- annealed and Optimized-annealed states.Fig. 42: Scanning Electron Micrographs of the partially annealed sandwich insert (expanded, to show middle core-shell fibersFig. 43: Effect of Annealing and E-Beam Sterilization of the Microstructure of the Insert (AB-NANOM-X-07-65).Fig. 44: Effect of Annealing on Release Rate of a Model protein (MW 85- 140K)Fig. 45: Effect of E-beam Sterilization on In-Vitro Release of Incorporated Protein (MW <25K)DETAILED DESCRIPTION

[0020] The invention describes several different strategies to deliver therapeutic molecules (proteins, peptides, nucleic acids, or small molecules) through electrospun (a) “layered” (sandwich) inserts, while also providing means of modulating release of bioactive substance, protecting the integrity of encapsulated substance, increasing biocompatibility of said device with underlying tissue andretention of device at site of application, and (b) dissolvable inserts, with said insert dissolving on tissue surface in a thin transparent film after placement, forming an airoil-water interface on the tissue surface due to incorporated ingredients, with or without drug, for the purpose of alleviating a tissue surface disorder.

[0021] Sandwich (Layered) Inserts:

[0022] Compositions of fiber-based inserts produced by the process of electrospinning and electro- spraying described herein have been focused on (a) encapsulation of concentrated molecules with various levels of water solubility and lability (stability), (b) varying compositions and methods to release said molecules in sustained manner, (c) minimization of irritation at the local tissue, and (d) maximization of retention at the local site of delivery.

[0023] The inserts described, have a hydrophobic barrier to drug release, in the form of a hydrophobic outer "shell" layer encasing a hydrophilic "core" layer, which contains the drug. In order to accomplish a linear rate of release, a "sandwich" insert has been described, with the drug-contained fibers (core- shell, or monolithic) sandwiched by two polymeric layers ("Encasing Layers").

[0024] The sandwich devices have 3 layers. The 2 outer layers (encasing layers) have a hydrophobic polymer composition that modulates release of the encapsulated substance, provides flexibility, is tissue pliant (wraps on the tissue), has dimensional strength and stability upon hydration with tissue fluids. The two outer layers are electrospun and form fibers upon deposition, but "flow" upon drying to form a continuous, or almost-continuous film that covers / encases the inner layer.

[0025] The inner layer of the sandwich device consists of fibers (that remain as fibers, unlike the encasing layers which transform upon drying into the continuous encasing layers). The inner layer can contain a homo-fiber (single channel), or a coreshell fiber (dual channel), depending upon the release profile desired.

[0026] The encasing "layer" polymers have properties that allow them to "flow" to form the encasing layer. "Flow" can be obtained if the glass transition temperature of the encasing polymer

[0027] The encasing layer may be selected from the group that is derived from poly (trimethylene carbonate)(PTMC). The derivatives may be PTMC derivatized with polycaprolactone, poly(lactide), or poly(siloxane), to impart various properties to thefinal polymer. PTMC of one molecular weight can be blended with PTMC of another molecular weight, to provide flow properties to form a continuous, or almost continuous sandwich layer.

[0028] The purpose of generating a layered drug delivery system is to generate a flexible tissue-device interface to eliminate / minimize irritation using an elastomeric polymer. If the modulus of tissue and the modulus of insert is matched, or if the modulus of the insert < modulus of tissue, insert will not delaminate and slide on the tissue causing irritation. The properties of the encasing polymer layer enable tissue-conformance, which enhances retention on mucosal surfaces

[0029] The inner encased layer contains the drug that can be released in a sustained manner. To accomplish this, the drug may be contained within a core-shell fiber matrix. The drug optionally, may also be contained within the encasing layer.

[0030] The inserts may be inserts (thin films), annular ("ring"), elliptical, curved or tubular in shape. The tubes may be 1-10 mm is length, and 0.2-10 mm in diameter.

[0031] The drug in question, can be a macromolecule or a small molecule, or a combination of molecules. The molecule can be biologically derived, or synthesized. The molecule can be a protein, or part of a protein, a nucleic acid, or a peptide, or a peptide-nucleic acid, or a small molecule with varying degrees of water solubility and stability.

[0032] Due to the "thin" character of the insert, the said delivery device is biodegradable, and degrades by surface erosion.

[0033] All layers of the insert are fabricated by the process of electrospinning of fibers, which are deposited onto a collector. However, most polymers that are electrospun stay as fibers. To form the encasing layers, PTMC polymers, derivatives and blends of PTMC would need to be used. These types of layered drug delivery systems can be used to deliver water soluble compounds, water-insoluble compounds. These compounds can be small molecules, or macromolecular biologies such as antibodies, nucleic acids and inactivated viruses.

[0034] Attributes of the Inner Layer of the Sandwich Device

[0035] The drug delivery system (insert) has the drug encapsulated in an inner "core" (comprised of water-soluble, protective components) and encased by hydrophobic polymers to create a "shell", to achieve slow release from the matrix. For labile drug molecules in the "core", the polymer "shell" matrix is protective to degradative elements and is a barrier to drug release. The components of the core protect the drug from degradation (such as oxidation, proteolysis, deamidation).

[0036] An additional outer encasing layer may be deposited on the insert, "sandwiching" the core-shell fibers by encasing the fibers in a contiguous layer. The "sandwich" or encasing is to modulate drug release and to provide a flexible biocompatible surface that touches the tissue.

[0037] Encapsulation of Labile Water-Soluble Molecules in Inner "Core" of a sandwich device.

[0038] The drug delivery system (insert) has the drug encapsulated in a "core" by hydrophobic "shell" polymers, to achieve slow release from the matrix while being protected from endogenous proteases, nucleases or oxidative species that are known to trigger degradation. For protection of the drug, excipients can be included to protect the drug from oxidation, deamidation, aggregation.

[0039] Enabling Sustained Release of a Highly-Water Soluble Drug

[0040] Water-soluble bioactive molecules are extremely difficult to encapsulate in concentrations that enable long durations for drug release. One way to effectively encapsulate water-soluble proteins in a hydrophobic polymeric matrix is to generate inserts of the core- shell design, with the protein as the hydrophilic core and the annular polymer encasing as the hydrophobic barrier to release. In order to do the above, the water-soluble drug like a protein needs to be dissolved in an aqueous medium and the hydrophobic polymer dissolved in an organic solvent. Methods are disclosed below that minimize the precipitation of protein at the aqueous-organic interface. Methods of creating an interface that is biocompatible to ingredients present in both phases are also disclosed.

[0041] Another way to encapsulate drugs effectively is by "sandwiching" drug-containing electrospun fibers within two continuous film layers effectively sandwiching the drug into a fully encapsulated drug delivery system. The highlyelastomeric outer layers act to "mold" itself onto the tissue surface minimizing or eliminating any biomaterial-tissue reactions. Sandwiching the drug-containing electrospun fibers between two elastomeric hydrophobic membrane film layers that can enable sustained release over the requisite duration. The membrane sandwich layers are polymers that have flow properties once dried; polytrimethylene carbonate and derivatives thereof, have characteristics that enable the formation of membranes. The sandwich insert construct allows drug molecules to be encapsulated in the "middle" layer, while the outer layers act as barrier to release.

[0042] Use of Polymers with Dimensional and Modulus Stability

[0043] The selection of materials that do not become stiff (i.e., modulus stability), or contract (dimensional stability) as it sits in the biological milieu are key criteria of the invention. Certain biodegradable polymers (PLGA and PLGA copolymers) used in the insert become stiff and inflexible (change in modulus), and contract (change in dimensions) upon hydration in buffer at physiological temperature (37°C) and pH (7.2-7.4). This results in inserts turning into dimensionally diminished shards with sharp edges, resulting in delamination of the inserts from the tissue surface and irritation due to loss of flexibility. In contrast, the inserts constructed with blends of amorphous polymers such as polycaprolactone, polytrimethylene carbonate and poly (trimethylene carbonate: polycaprolactone), ethylene vinyl acetate polymers, polysiloxane and polydioxanone demonstrated dimensional stability, as well as sufficient elastomeric attributes to sit flexibly on the tissue surface. Polytrimethylene carbonate and polycaprolactone in a random co-polymer format in various ratios can be utilized to achieve inserts of varying strengths and elastomeric character. Blends of ethylene vinyl acetate (EVA) with polycaprolactone demonstrated dimensional stability and elastomeric qualities. Copolymers or blends of various combinations polytrimethylene carbonate, polycaprolactone, polydioxanone, ethylene- vinyl acetate, polysiloxane are appropriate for this purpose.

[0044] The invention describes compositions and methods to formulate drugcontaining inserts fabricated by the process of electrospinning, electrospraying, or a combination of electrospinning and electrospraying, for the purpose of applying them to tissue surfaces or implanting them within tissue. The purpose of the invention is to enhance biocompatibility of the biomaterial-tissue interface by minimization ofirritation, enhance retention of the insert within the tissue by matching flexural and compressive modulus of tissue to polymeric insert material and modulate release of the encapsulated drug, by modulation of matrix polymer composition and microstructure of the encapsulating polymer layer.

[0045] The insert can be placed within a tissue space as an implant, whereupon the implant is surgically placed into the tissue space. For subcutaneous space, the insert containing drug can be placed surgically in the tissue space. The insert can also be implanted into tissue that can have large amounts of liquid such as the urological space, as in the bladder. In another embodiment, the topical insert can be placed in sinusoidal cavities. In another embodiment, a sustained drug-eluting insert can be placed in other tissue surfaces, such as rectal tissue, vaginal tissue and esophageal tissue. For other tissues that can benefit from drug delivery inserts, such as intranasal, intra-bladder, sub-lingual, intra-gingival, buccal, intra-vaginal, intra- esophageal, intra-brain, intra-rectal tissues, identical or similar considerations apply.

[0046] Once adhered to the tissue, an insert that is bonded to the tissue (by ionic bonding, or hydrogen-bonding) can be removed by application of an isotonic salt / excipient composition. Compositions of hydrating solutions are disclosed that will de-adhere the insert from mucosal tissue surfaces.

[0047] In one embodiment, the insert design incorporates a hydrogel encasement on each insert fiber that adheres to the tissue surface, while also enhancing biocompatibility due to its high-water content. The inner insert fiber layer contains the drug, in the "core". The drug may be hydrophobic, hydrophilic or amphiphilic. The hydrogel encasement is the "sheath" encapsulating each fiber.

[0048] The adherence of the insert is enhanced by the application of a hydrating fluid (wetting solution), with adherence within 1-30 minutes.

[0049] In another embodiment, the insert form incorporates an inner hydrophilic core (that also contains the drug) and an outer hydrophobic flexible, elastomeric sheath, and co-spun with a biocompatible adhesive. The inserts are removable with the application of a fluid, and applied with a fit-for-purpose designed applicator. For The inserts are sterilized under certain sterilizing conditions. In another embodiment, the drug delivery system does not have a bioadhesive, but is held-in-place by tissue geometry, and / or tissue folds. Additionally, methods andcompositions are disclosed to encapsulate and sustain-release intact, highly water soluble, bioactive compounds such as proteins, peptides, nucleic acids and small molecules in electrospun nanofibrous inserts.

[0050] Enhancing Biocompatibility and Retention by Encasement of Each Fiber of the Insert.

[0051] One strategy to improve the issue of biocompatibility and retention is by generation of a nano-porous mesh insert that contains the hydrophobic biodegradable fibers individually encased in hydrophilic, mutually- reactive and tissue-reactive polymers that reacts with one other and with the tissue proteins forming amide covalent linkages. This tethers the drug delivery system to the tissue site and encases the hydrophobic polymer fibers. In the case of PLGA, which degrades by bulk erosion, encasement by a hydrophilic sheath layer prevents formation of the particulate debris in the local tissue space, ensuring long-term biocompatibility of the drug delivery system. The composition of the encasement is focused on polyethylene oxide branched polymers with reactive functionalities that are designed not only to react with one another, but also with tissue proteins. The presence of polyethylene oxide encourages additional biocompatibility, due to high water content. Thus, the immunomodulatory cells of the body "see" a highly hydrophilic surface and neutral pH. The encasement also contains polyvinyl alcohol, a filler molecule that allows the modulation of crosslink density of the hydrophilic annular sheath.

[0052] The second type of electrospun fiber-based insert composition is a dissolvable insert, which placed on tissue would transform into a hydrogel and optionally, crosslink with itself and with tissue and slowly dissolve over time in the hydrating tissue milieu, either as a means to lubricate the eye, or release a compound which may be a drug molecule, as the insert dissolves. The dissolvable insert may start out as an opaque insert, but slowly turn transparent or translucent as it absorbs water. The dissolution time would range from immediate after placement to 24 hours after placement.

[0053] The term “dissolvable” implies slow, or fast dissolution of the insert in the tissue space that this is placed in. All components in the dissolvable insert dissolve by dissolution or emulsification of the components. The inserts are dissolvable, tissue-conforming, highly hydrating and adherent to tissue. Dissolvable inserts can be used in all mucosal tissues.

[0054] The foci of the drug-delivery system were to: (a) generate a biocompatible, highly flexible electrospun surface, so the biomaterial-tissue interface generates minimal or no irritation, (b) generate insert formats that will achieve sustained release of proteins, nucleic acids and small molecules, (c) generate insert formats that encapsulates high concentrations of proteins without causing the generation of aggregated proteins, (d) generate insert compositions that can be sterilized by various sterilizing means, and yet not result in loss of integrity of the encapsulated protein.

[0055] The "sandwich" insert has three layers: (a) a lower hydrophobic, elastomeric layer, (b) a middle layer containing the therapeutic to be delivered and (c) upper hydrophobic, elastomeric layer. Other variations of the "sandwich" insert would be a simpler middle layer with any polymer with any drug incorporated, be it hydrophilic (MW>100), or hydrophobic (MW>100), synthetic or biologic. Thus, the middle layer could contain any drug class. The properties of the polymers that comprise the insert may not be ionic, or charged, but elastomeric to create Van Der Waal type forces with the tissue surface that results in long term adherence.

[0056] The tissue types may be dermal, ocular, vaginal, intranasal, rectal, esophageal, urological, esophageal, sublingual, or buccal. The shape of the device thus formed can be in the form of a round-shaped disc, an oval-shaped disc, a shape that is contoured specifically to fit in the cul-de-sac of the conjunctiva, a ring-shape that fits on the ocular surface. For ocular use, the round-shape disc may be 2 mm to 10 mm in diameter, preferably 3 mm to 6 mm in diameter, most preferably 4 mm- 6 mm in diameter. For dermal use, the device diameter can be 10-50 mm in diameter. For intranasal use, the device diameter is less than 10 mm in diameter. For urological use, the device may be inserted into the bladder with the assistance of a catheter. The device for urological drug delivery may be tubular in structure.

[0057] Attributes of the Encasing Layer

[0058] The top and bottom layer would contain an elastomeric polymer such as poly(TMC), poly(TMC-polycaprolactone), poly(caprolactone), polydioxanone, ethylene-vinyl acetate and blends thereof to achieve the properties that would lead toa continuous film layer generated by electrospinning. In one variation, the top and bottom layer would contain a PEGylated compound to modulate release of drug from the middle layer (diffusion aid).

[0059] Inserts have Elastomeric Elongation >100%

[0060] The inserts have elastomeric elongation > 100%, to be conforming to tissue. To test the mechanical adhesive strength of inserts to tissue, a stress-strain curve was generated to determine the strain at break. Fig. 27 demonstrates a >100% elongation of the insert. The testing was performed using a Shimadzu Mechanical Tester.

[0061] An experiment was performed with fresh bovine conjunctiva to compare inserts with different elastomeric elongation percentages. Inserts having elongation (% strain) < 100%, were not pliable and did not conform to the underlying tissue. In contrast, inserts with elongation > 100% conformed to the underlying tissue.

[0062] Attributes of the Encased (Middle) Layer

[0063] The middle layer has a core-sheath structure, and optionally, a monolithic layer (not core- sheath).

[0064] The purpose of the middle layer to encapsulate the hydrophilic compound so that: (a) high concentrations of the compound (protein, etc.) can be encapsulated in a stable manner, (b) excipients that stabilize the protein can stay encapsulated with the protein. The water-soluble drug is incorporated within the core of an annular fiber structure of the electrospun insert, with the sheath containing the encapsulating hydrophobic polymer. The capacity of the sheath to function as a contiguous layer defines the ability to slow down release of an encapsulated compound, although the top / bottom layers of the insert would primarily perform that function.

[0065] The microstructure of the insert may be such that labile water-soluble molecules are encapsulated within core-shell fibers of the insert, to enable protection from proteases, nucleases other enzymes and oxidative degradation. The drug delivery system (insert) has the drug encapsulated in a core by hydrophobic "shell" polymers, to slow down release from the matrix while being protected from endogenous proteases, nucleases or oxidative species that are known to trigger degradation.

[0066] For protection of the drug, excipients can be included to protect the drug from oxidation, deamidation, etc. The outer shell is hydrophobic and is the encapsulating polymer, ideally causing a continuous layer. The drug molecules encapsulated in the "core" can be hydrophilic, hydrophobic, or amphiphilic. The drug molecules could also be hydrophobic as a free base or the free acid, but rendered water-soluble by conversion into a salt form.

[0067] The salt forms can be selected from salt form available in the list of pharmaceutical salt form, including but not limited to, acetate, hydrochloride, malate, iodide, bromide, chloride, malonate, tartrate, gluconate, glycolate, or any other salt form that may render a water-insoluble drug soluble in water. The active molecules may be amphiphilic, i.e., soluble in both aqueous and non-aqueous media. The active molecules may be selectively soluble as a function of pH, with a pH range of 4.2 to 7.4, and a preferred pH range of 5.5-6.5 and another preferred pH range 4.2-5.5.

[0068] The active drug molecule may be a biologic in nature, such as an antioxidant, peptide, a peptide-nucleic acid, plasmid DNA or linear DNA, a protein. The drug may be biologically derived or synthetic, or a synthetic derivative from a biological molecule. The preferred molecule is a protein, with a preferred molecular weight between 10-30 KD, a preferred molecular wright between 30kD and 100KD, a preferred molecular weight between 100KD and 150KD and a preferred molecular weight between 150-250KD. The molecule may be a nucleic acid, such as an antisense oligonucleotide, or messenger RNA, or RNAi, or aptamer, or a peptidenucleic acid (PNA). The molecule may be either small molecule or macromolecule, a water-soluble molecule, or a molecule rendered less water soluble by salt formation, or derivatization with another molecule. The molecule may be anti-glaucoma, antiinflammatory, anti-microbial. The molecule may be a muscarinic agent, or an agent that has a therapeutic effect. The active may be complexed with another molecule to enable higher encapsulation. The active drug contained in the "core" may be complexed into nanoforms such as calcium complexes of DNA, positively charged peptide complexes with negatively charged polymers such as hyaluronate, xanthan.

[0069] The drug-containing solution may contain dissolved excipients that lower the surface tension, to enable the process of electrospinning. The surface tension of the aqueous liquid (which contains the water-soluble active ingredient) is inthe range 25-55 mN / m, measured by a Attension™ Theta Contact Angle Optical Tensiometer. Preferably, surface tension of the aqueous liquid is in the range 20-45 mN / m. In a more preferable range, the surface tension of the aqueous drugcontaining core is in the range 30-40 mN / m, and most preferably, in the range 30-36 mN / m. The conductivity of the aqueous core is in the range 2000-6000 pS / cm, with a preferable range between 2000-4000 pS / cm. The viscosity of the aqueous core solution containing the drug (peptide, protein, small molecules, nucleic acids) is in the range 1-900 cP (25°C, shear rate 40 s-1), viscosity of the encapsulating layer is in the range 1500-10,000 cP, with a preferable viscosity range as 1500-6000 cP, measured with an Anton Paar MCR92 Cone-Plate Rheometer using a spindle CP50-0.5 / T. The interfacial tension between the aqueous and the organic phase is less than 30 mN / m, preferably <20 mN / m. Electrospinning of the core solution and encapsulating polymer solution is performed with a tri-axial set-up, with an organic vapor flowing at a low flow rate of 70 mL / min. The organic vapor is a class III solvent and preferably methyl acetate, or acetone. The distance from the emitter (needle tip) to the collector is in the range 15-21 cm, with a preferable range 16-19 cm. The temperature of the process chamber is in the range 15-25°C, and relative humidity (RH) in the process chamber in the range 25-40% RH, with a preferable range less than between 25-35% RH, and a most preferable range between 25-30% RH. The encapsulating polymer solution can be at room temperature or slightly heated, within an acceptable range of 25-50°C.

[0070] The aqueous core-solution containing the dissolved water-soluble drug may contain excipients that are protective to the active ingredient. These excipients include but are not limited to histidine, glycine, methionine, sorbitol, sucrose, trehalose, sodium chloride, glycerin, hydroxypropyl beta cyclodextrin, cyclodextrins, polyethylene glycol 400, polyethylene glycol 8000-200,000, propylene glycol, polypropylene glycol 400-5,000, polyvinyl alcohol, polyvinyl pyrrolidone, PEG2K- DSPE, PEG5K-DSPE, PEG2K-DPPC, PEGylated phosphatidyl choline, phosphatidic acid, lecithin. The aqueous solution may contain a protein as an excipient, such as collagen, or albumin.

[0071] Use of Excipients to Generate an Interface between Aqueous and Organic Phases

[0072] Proteins are typically unstable at aqueous-organic interfaces, where denaturation and loss of tertiary structure is common. The working hypothesis was that having a monolayer of an amphiphilic surfactant at the interface would serve to protect the protein from immediate events of denaturation resulting in irreversible aggregation and precipitation. We hypothesized that a molecular interfacial layer between the aqueous and organic phases would proffer protection to the protein from having prolonged contact with the organic phase. We further hypothesized that PEG2KDSPE would remain at the interface since it is soluble in both layers, would lower the surface tension of the aqueous phase allowing the protein to be electrospun into core-sheath filaments, encapsulated by the sheath polymer. Previously, the protein could not be electrospun due to instant precipitation and aggregation, and an increased interfacial tension at the core: sheath interface when the two solutions make contact. Experiment: The experiment evaluated if an immunoglobulin protein at 20 mg / g could be protected from denaturation by the inclusion of this excipient in the core. Measures of success were: (a) visual assessment of the interfacial layer and assessment of lack, or presence of precipitation of protein, (b) assessment of protein integrity by assessment of % recovery from expected value, assessment of % HMW, in relation to percent Monomer by Size Exclusion Chromatography (SEC). The aqueous "core" solutions all contained protein at 20 mg / g, reconstituted first at 50 mg / g with a phosphate buffer solution. Data shows that increased concentration of PEG 2k DSPE above 0.1% did not provide any additional protection at the core (aqueous) / sheath (organic) interface, surface tension results increased to > 41 mN / m as PEG 2k DSPE increased, indicating that above 0.1% = CMC (critical micelle concentration).

[0073] Concentrations: Solutions containing 0.25% and 0.4% PEG 2k DSPE, with and without methyl acetate, resulted in low protein recovery, 60 - 79% at T=0, with lower recoveries at T=2 hours. Solutions containing ?0.1 % PEG 2k DSPE, with and without methyl acetate, provided protection at the interface of the two phases and protein recovery varied from 81 - 103%, staying fairly consistent across the board for T=0 and T=2 hours.

[0074] The interface between aqueous and the organic phase is addressed by the presence of PEG2K-DSPE in the aqueous phase and span 40 in the organic phase.The concentration of PEG-DSPE can be 0.05% to 1% in the aqueous phase, with a preferable concentration range of 0.1% to 0.4%. The PEG-DSPE may be PEG2K- DSPE or PEG5K-DSPE.

[0075] The (shell) polymer may be biodegradable or durable. The encapsulating polymer is dissolved in an organic solvent such as dichloromethane (DCM), methyl acetate, ethyl acetate, acetone, butanone, or tetrahydrofuran, mixtures and blends thereof, or a suitable organic solvent that dissolves the polymer and does not have a large diffusion coefficient into aqueous. The concentration of the polymer dissolved in the organic solvent is in the range 5-50%, depending upon the molecular weight of the polymer. The preferred range is 10-20%.

[0076] The shell polymer may also be dissolved with an acetic acid component to increase polarity of the solution, provided the drug is compatible with acetic acid. For example, a small molecule drug may be compatible with acetic acid in the formulation, but a macromolecule such as protein or DNA will not be. For drugs that are stable at a pH < 5, acetic acid may be a good

[0077] The encapsulating (shell) polymer may be ethylene- vinyl acetate (EVA), with 25-45% vinyl acetate content, with 38-42% being the preferred range. The shell polymer may be a blend of EVA and other polymer to provide the resultant blend with desired biomaterial properties. The shell polymer may also be polytrimethylene carbonate, or polytrimethylene carbonate-co-polycaprolactone, or polytrimethylene carbonate-co-DLL-lactide used as a single polymer or in suitable blends with other polymers either hydrophobic or hydrophilic, either to render strength, enhancement of elastomeric properties and creation of a contiguous encapsulating layer with no pores or breaks. Other polymers may be poly anhydrides, polyester- amides, polycaprolactones, polycaprolactone-TMC, polylactides, PLGA and blends, or copolymers thereof.

[0078] Incorporating Very High Concentrations of Protein in the Insert

[0079] The protein is contained in the middle layer of the "sandwich" insert. The types of proteins are immunoglobulins, globular proteins, recombinant proteins, fusion proteins, chimeric proteins, antibodies, structural proteins, growth factors, enzymes. A stable formulation is one where protein stability is maintained through its physical stability, chemical stability and biological activity. Physical stability can bedetected either in the formation of precipitates (particles) or high MW aggregates by Size Exclusion Chromatography. Chemical degradation can be size clipping, deamidation, methionine oxidation, disulfide scrambling, or other changes to the primary structure of the protein. Both physical and chemical changes can lower the biological potency of the protein.

[0080] One of the key challenges for the development of sustained release protein-containing dose forms is its susceptibility to its microenvironment, often governed by its concentration. It is difficult to encapsulate high concentrations of proteins in a polymeric matrix. Upon hydration of the dried insert, proteins in the aqueous microenvironment (such as tear fluid) are at a very high concentration (>100 mg / mL), which increases opportunities for aggregation, both irreversible and reversible. Stabilization of proteins and other biologic drugs is crucial for ensuring both safety and efficacy. Molecular structures of therapeutic proteins and the critical sites within these proteins that are prone to oxidation, deamidation, hydrolysis, etc., and can result in aggregation and fragmentation.

[0081] The development process for a protein-based drug product, therefore, begins with efforts to stabilize therapeutic proteins. The unstable nature of proteins make stabilization challenging; the folded state is only marginally more stable than the unfolded state, any changes in the protein environment may trigger protein degradation, aggregation, and or inactivation.

[0082] Temperature and pH are parameters that can affect stability of protein, leading to thermal denaturation and aggregation, leading to chemical degradation such as side-chain oxidation, hydrolysis and deamidation. For any particular protein, identifying a select window of pH for maximum protein stability is crucial. The effect of pH on the chemical stability of a protein can be altered in the presence of excipients. The effect of salts strongly depends on the pH of the solution, which influences the charged state of ionizable groups in the protein. Depending on the type and concentration, metal ions may also stabilize or destabilize a protein because their interactions with proteins are highly protein-dependent. Metal ions may also significantly affect protein stability without affecting much of its secondary structure. These target the Fenton pathway, targeting methionine, cysteine, histidine, tryptophan, tyrosine, proline, arginine, lysine, or threonine. The catalysis depends onthe concentration of the metal ions, and the metal-catalyzed reaction can be facilitated in the presence of reducing agents such as ascorbate or sulfhydryl (RSH) compounds.

[0083] Chelating agents such as ethylenediaminetetraacetic acid (EDTA) and citric acid can bind and destabilize proteins or bind harmful metal ions and prevent oxidation. Protein concentration can also influence aggregation and in some cases chemical degradation, depending on the mechanism. Higher concentrations create more chances for unwanted self-association, and consequently high viscosities.

[0084] Chemical degradation pathways include not only oxidation, hydrolysis, and deamidation, but also isomerization, succinimidation, disulfide bond formation and breakage, non-disulfide crosslinking, and deglycosylation. Deamidation of asparagine and glutamine residues is most common, and the rate, mechanism, and location of deamidation are all pH-dependent. Oxidation is another important chemical degradation mechanism, particularly the oxidation of the thio groups in methionine and cysteine residues. The formation of disulfide bond linkages or thiodisulfide exchanges can result in protein aggregation or polymerization.

[0085] Dissolvable Inserts

[0086] Topical administration to the eye involves eye-drops, which rapidly clear from the surface and result in low bioavailability (<10%) due to drainage from naso-lacrimal channel, overflow due to induction of tearing and blinking. To minimize this effect, a slow-dissolving insert can efficiently release polymer excipients and / or drug, without overwhelming the ocular space.

[0087] Disclosed here are variants of water-dissolvable inserts with both water soluble and oil excipients incorporated, applied to the tissue surface as dry polymeric inserts, which upon hydration absorb water to transform into a hydrogel, conforming to underlying tissue, optionally crosslinking onto the tissue. The transparent hydrogel forms an air-oil-water interfacial film on the hydrated tissue surface, thus preventing evaporation of water from the tissue surface. Hydrogel dissolves away slowly based on its composition and its dissolution time is modulated by composition of the insert. The purpose of an oil-water insert is to form a "monolayer" interface on the tissue surface, with the oil floating to the top, preventing evaporation.

[0088] In another variant, two different fibers are electrospun to form an insert construct. When in contact with water, components in one fiber can react withcomponents in the other fiber to form an in-situ forming hydrogel film on the tissue surface, that dissolves / biodegrades slowly in the biological milieu.

[0089] The water-dissolvable inserts may act as a bandage (with no drug) to allow a wound to heal, or optionally contain a pharmaceutical drug with a therapeutic effect, or a biologic with biological pharmacology, or a polymer, or polymers that are co-electrospun to mutually crosslink with one another and the tissue. The dissolution time can be varied (1 hour to 2 weeks) by varying the type of polymer included in the electrospun and its concentration, thereof. The water-dissolvable inserts can be applied onto tissue surfaces. The type of tissue is ocular, intranasal, vaginal, rectal and oral. The water-dissolvable inserts may be applied on a wound, or abrasion as a bandage barrier, optionally containing an anti-microbial, pain reliever or a drug that is a wound healing compound. An example of this is a corneal bandage, for wound healing of corneal abrasions. The benefit of this is due to the transparent nature of the film, there is no vision impairment due to the device. Another example is an oral mucosa-adherent insert to cover canker sores, while lubricating the oral cavity. An example is an oral insert containing ingredients to alleviate dry mouth. Another example is an intravaginal insert to alleviate dryness of the vaginal mucosa. Another example of utility for a dissolvable insert is one that is incorporated with clotting factors and that can be packed into a deep wound to stop bleeding.

[0090] Attributes of the Dissolvable Insert

[0091] Dissolution time of the insert can be an hour, or it can be 12-24 hours, or even 1 day to 14 days. A water-soluble insert can efficiently deliver ingredients to the ocular surface, or any mucosal surface. The ingredients may comprise a drug or polymeric excipients, being released to the ocular surface. The ingredients are a mixture of oils and aqueous polymers, emulsified and electrospun to form an air-oilwater interface. Inserts are normally designed to be placed under the lower eyelid in the conjunctival fornix, but a fast-dissolving thin hydrogel may be placed on the sclera (white part of the eye) to enable better absorption of the ingredient. For treatment of ocular surface disorders, it is logical to place the insert on the corneal surface, as in Fig. 3. The ocular insert may be placed in the conjunctival fornix (lower eyelid), for ease of placement and fast dissolution.

[0092] Variants of water-dissolvable, biodegradable insert were developed and formulated with ingredients that are biocompatible and soothing to the tissue surface. The inserts are of the shape circular, or a shape that is curved to fit the contours of the underlying tissue. The insert can be placed on the cornea to aid healing, or to aid in a surgical procedure. The insert as a dry device is white and opaque, but after hydration can be transparent, or translucent. The inserts can be sterilized by e beam sterilization at doses between 6-25 kGy, and gamma irradiation at doses between 6-25 kGy, cold.

[0093] The composition of the insert may be varied to vary the dissolution time of the insert in aqueous fluid of the eye. This can be accomplished by incorporation of water-soluble polymers of different molecular weights. These can be selected from, e.g., polyethylene glycol, polyvinyl alcohol, hyaluronate, sodium alginate, xanthan, pectin, celluloses (carboxymethyl-, hydroxypropyl-, hydroxypropylmethyl-), guar gum, tamarind seed polysaccharide, l->-3 Beta Glucan, poly(arginine), and polylysine to treat ocular dryness and lack of lubrication and ocular itching due to allergens or dryness. These inserts are dissolvable instantly, in less than an hour or prolonged, with dissolution time modulated by insert composition. Other tissue dryness can be alleviated as well, such as oral dryness and vaginal discomfort and dryness. The insert is comprised of one or more mucoadhesive polymers, and one or more biocompatible polymers. Biocompatible electrospun insert is comprised of two different compositions of fibers, with components in one type of fiber optionally capable of reaction with components included in the other type of fiber, both co-spun simultaneously by electrospinning, to form a water-dissolvable nano structured insert. The structure of the insert is nano structured, generated by the process of electrospinning, or electrospraying, or a combination of the two processes. The nanostructure allows the insert to rapidly absorb water, optionally mutually crosslink and with the underlying tissue. By virtue of light crosslinking, the insert does not immediately "ball" up into gel pieces, but instead maintains its shape. The dry device will dissolve in water and water-containing medium, including biological media of the type tear fluid, nasal fluid, oral fluid, vaginal fluid and rectal fluid.

[0094] The dry insert product can be produced by the electrospinning of fibers. Two different types of fibers can be co-spun by the process of electrospinningforming a single insert with the first fiber containing components optionally reacting with components contained in the other type of fiber. In one example, one type of fiber contains sodium hyaluronate of MW in the range 50,000-5000,000 Daltons, polysorbate 20 or polysorbate 80 and polyvinyl alcohol of MW in the range 20- 130,000 Daltons and optionally, sodium chloride in the range 0.1-3.5%, boric acid or its sodium salt thereof in the range 0.0001% to 5%, sodium alginate in the range 0.01%-20%, polyethylene glycol with MW in the range 100-100,000 Daltons and between 0.01-60%, dextran 70 in the range 0.01-50%, calcium chloride in the range 0.0001%-l%, polyvinyl pyrrolidone in the range 0.1-50%, albumin in the range 0.01- 10% and the second fiber contained in the insert is comprised of polyvinyl alcohol of MW in the range 20-130,000 Daltons, and, optionally, calcium chloride (0.0001-1%), sodium borate (0.0001-1%), dextran 70 (1-50%), polyethylene glycol (of MW between 100-100,000 Daltons), PEG-multi-arm-succinimidyl ester (with the number of arms between 4-16), polypropylene glycol (MW 400-4000), propylene glycol, collagen type I, glycerol, a bioactive substance, sodium alginate in the range 0.001- 20%, sodium hyaluronate (MW in the range 50,000-5000,000 Daltons and compositionally between 0.01-1%, xanthan gum (0.01-10%), guar gum (0.01-10%), hydroxypropyl beta cyclodextrin, trehalose, mannitol, sorbitol, glycerol, tamarind seed polysaccharide, or beta 1>6 glucan, wherein the first fiber and the second fiber are configured to react with each other and with a biological tissue upon hydration, and the dry device transform into a hydrogel upon contact with the biological tissue and subsequent hydration and completely dissolves over a period of 0 h to 14 days after contacting the biological tissue.

[0095] The duration of dissolution can be varied by altering the concentrations of the cross -linkable moieties such as alginate crosslinking with calcium chloride, or another type of divalent salt such as magnesium chloride, or zinc chloride. Another type cross -linkable moiety is boric acid crosslinking with polyvinyl alcohol or other multi-hydroxyl containing species such as glycerol, sorbitol, or mannitol. In another embodiment of this concept, polyethylene oxide multi-arm amine can be varied in concentration along with polyethylene oxide multi-arm succinimidyl ester to form a loose or tightly crosslinked hydrogel. It is not necessary that each type of fiber actually be crosslinked tightly; optionally, each type of fiber may contain polymersthat slow dissolve and may not be crosslinked at all, should the dry device be required to dissolve quickly, within 2 hours of placement onto the tissue.

[0096] For electrospinning, the solutions into fibers, the viscosity of PVA based solutions is in the range 1000-5000 centipoise at 25°C, viscosity of PEG solutions is in the range 7000-15000 centipoise; surface tension is between 25-50 mN / cm and conductivity is between 1-30 mS / cm.

[0097] In another example, the insert is comprised of hyaluronate and polyvinyl alcohol in the first fiber and polyethylene glycol and alginate in the other fiber type. Boric acid and calcium chloride can be used in the fiber compositions to crosslink the insert upon hydration.

[0098] The insert is administered as a dry device, which rapidly absorbs water from the surface of the eye, or the soft tissue on which it is placed. The insert upon hydration, absorbs water from the surrounding tissue to form a tissue adherent hydrogel.

[0099] The insert device comprised of electrospun fibers may optionally, cross-react with components in the device and including water, transforms into a tissue-compliant hydrogel. Optionally, the insert may not have components that react with one another. The two fibers may have the same components or different components.

[0100] The hydrogel dissolves over time, with dissolution time modulated by composition. Upon dissolution, the hydrogel releases its ingredients slowly over time. By slow release of high viscosity polymeric components, the ocular space is not overwhelmed as in eye-drops which induces higher volumes of tear fluid and tear flow.

[0101] The dry insert can contain tissue-lubricating excipients including but not limited to hyaluronate (0.05-15%, MW 80,000-3,000000 Daltons), xanthan (0.05- 15%), guar (0.05-15%), gelatin (0.05-15%), soluble collagen (0.1-17%), hydroxypropyl cellulose (HPC)(0.1-15%), hydroxypropyl methyl cellulose (0.05-5%) lecithin (0.001-1%), dextran (0.05-15%), carboxymethyl cellulose (0.05-15%), glycerin (0.05-25%), tamarind seed polysaccharide (TSP)(0.05-15%), polyvinyl alcohol (MW 20,000 Daltons to 130,000 Daltons), polyethylene glycol 100-100,000, polyethylene glycol 100, polyethylene glycol 8000, polyethylene glycol 20K,polyethylene glycol 35K, trehalose, mannitol, glycerol, propylene glycol, polypropylene glycol 400-4000, PEG40-stearate, poloxamer, polyvinyl pyrrolidone (PVP), tamarind seed polysaccharide (TSP), 1-3 beta glucan, pectin, polyacrylic acid, starch, chitosan gluconate, albumin, beta glucan, gelatin, collagen, atelocollagen, lecithin, pegylated lipids, distearoyl phosphatidyl choline, dimyristoyl phosphatidyl choline, dimyristoyl / distearoyl phosphatidyl glycerol, hydroxypropyl methyl cellulose, hydroxypropyl cellulose and mixtures thereof, in concentrations from 0.01- 1%, 1-10%, 10-15%.

[0102] The insert can contain water-insoluble excipients that are beneficial to tissue, including oils such as castor oil, peanut oil, olive oil, medium chain triglycerides, com oil, oil of eugenol, sunflower oil, safflower oil, flax seed oil, triglycerides and diglycerides, menthol, camphor, ceramide, phospholipids, including derivatives and blends and in concentrations 0.01-30% in the final insert product. The benefit to having water-soluble excipients incorporated in a water-soluble polymeric solution, is to create an air-oil-water interface to prevent evaporation of water from the tissue surface. This is particularly useful in the treatment of dry eye. This is also useful in treating dryness of other tissue, such as vaginal dryness, and dry mouth. The insert may be useful in preventing wounds from drying out, as drying of wounds aids in car formation.

[0103] The aqueous-dissolvable insert may optionally contain compounds that include, but are not limited to, brimonidine, brinzolamide, timolol, prostaglandins, MMP-modulators, anti-oxidants, olapatadine, antifungals, naphazoline hydrochloride, epinephrine, other vasoconstrictors, melatonin, lipoic acid, Epitalon, n-acetyl cysteine, glutathione, cysteamine, superoxide dismutase, SOD mimetics, aldehyde traps, vascular endothelial growth factor traps, immunoglobulins, RNA, DNA, peptide- DNA constructs, clotting factors, FGF, NGF, EDGE, anti-VEGF molecules, growth factors, cell-penetrating peptides, polyarginine, polylysine, albumins, heparins, n- acetyl cysteine, amino acids, thymosin beta, plant flavonoids, antibodies, rock inhibitors, mast cell inhibitors, calcineurin inhibitors, PDGF-inhibitors, analgesics, anti-glaucoma agents, anti-histamines, anti-allergenics, vasoconstrictors, UV-block compounds, polyphenolics, wound healing agents, ROS quenching agents, redness relievers, lubricants, NO-releasing compounds, estrogens, estrogen mimetics, anti-bacterials, anti-virals, anti-fungal, anti-emetic, anti-pain, NSAIDs, analgesics, cannabinoids, vitamins, oils, triglycerides, cholesterols, vitamin E, vitamin D, vitamin B12, DNA including ministering DNA, plasmid DNA, RNA and CRISPR CAS9 constructs. . The dissolvable inserts can be utilized as an artificial lubricant, as a slow dissolving resource for lubrication in the eye, or as a surgical aid, or as a vehicle to release a molecule to relieve redness, or irritation, or an emollient. The dissolvable insert can be utilized to release a protein, DNA, mRNA, a RNAi, a CRISPR construct, or a small molecule < 1000 Daltons. A medication can be incorporated into the insert, whereupon the medication can be complexed prior to incorporation, or not complexed prior to incorporation.

[0104] The insert is produced by the process of electrospinning of each fiber type, by simultaneously spinning them (co-spinning) and collecting on a collector, thus producing a mixture of fibers. Each fiber type may be identical in composition, slightly different with slight modifications, or completely different. Once hydrated, the insert transforms into a tissue-compliant hydrogel, with tensile strength and Young's modulus < modulus of soft tissue.

[0105] The Young's modulus is very important to consider as a biomaterial that is stiffer than the applied tissue will cause irritation and if it is not stiff enough, it will not be able to maintain its structure.

[0106] For soft tissues and organs, the range of Young's modulus is between 0.0001 MPa to 1 MPa . Specifically, the brain has a Young's modulus of 0.0001- 0.001 MPa, skin has a modulus of between 0.001-0.01 MPa, the spleen and pancreas have modulus values in the range of 0.0025-0.005 MPa, and glands and muscles have modulus values of between 0.008-0.017 MPa.

[0107] The liver has a Young's modulus of about 0.0105 MPa and human lateral nasal cartilage has a modulus of 0.98 MPa. Additionally, data for the ultimate tensile strength of soft tissue is also available. The conjunctival fornix has an ultimate tensile strength of 0.7 MPa and a Young's modulus of 3.9 MPa, and the amniotic membrane has an ultimate tensile strength of 1.7 MPa and a Young's modulus of 11.5 MPa.

[0108] As shown in Fig. 4, the Young's Modulus of all dry inserts was less than 0.5 MPa.

[0109] The Young's Modulus of hydrated insert inserts were in similar ranges as soft tissue moduli, i.e., <0.5 MPa (Fig. 5). This allows seamless integration of biomaterial with underlying tissue.

[0110] The internal microstructure of the insert is comprised of two different fibers. The straight fibers are PEG-based, and the fibers with curving loops are alginate based (Fig. 6). The dissolvable inserts can be sterilized by e-beam, cold, or by gamma irradiation, cold.

[0111] The invention encompasses the following aspects:

[0112] Aspect 1: A dry drug delivery system comprising a middle layer encased in two outer layers, the middle layer and the two outer layers each comprising a mixture of electrospun fibers, the dry drug delivery system having a final combined density in the range 0.1-1 g / cm3, a thickness < 1 mm, an elastic modulus < 4 MPa, and a change in elastomeric strain >100%, wherein the two outer layers each include one or more polymers selected from ethylene vinyl acetate, poly trimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) having a copolymer ratio of 90:10, PCL, and P(TMC) -polyethylene oxide (PEG)-P(TMC) in which a molecular weight of the PEG segment is 1,000-20,000 Dal and a molecular weight of the P(TMC) segment is 1,000-10,000 Dal; and the middle layer being formed of coreshell fibers, each of which includes: a core that contains a bioactive substance and one or more excipients to protect the bioactive substance from degradation, and an outer shell formed of a hydrophobic elastomeric polymer comprised of P(TMC) and PCL in a weight ratio of 90:10 to 30:70.

[0113] Aspect 2: The dry drug delivery system of Aspect 1, wherein the ethylene vinyl acetate has a degree of vinyl acetate substitution of 42%, the P(TMC) has an inherent viscosity of 0.3- 1.2 dL / g, the P(TMC)-co-PCL has an inherent viscosity of 1.2- 1.6 dL / g) and a copolymer ratio of 90:10, and the PCL has an inherent viscosity of 1-1.5 dL / g). wherein the mixture of electrospun fibers in the two outer layers are annealed to form a continuous layer.

[0114] Aspect 3: The dry drug delivery system of Aspect 1 or Aspect 2, wherein the mixture of electrospun fibers in the two outer layers are annealed to form a continuous layer.

[0115] Aspect 4: The dry drug delivery system of any one of Aspect 1 to 3, wherein a content of the one or more polymers in the two outer layers is 0-10% w / w P(TMC), 60-95% w / w P(TMC)-co-PCL, 5-20% w / w PCL, and 0-2.5% P(TMC)- PEG-P(TMC), based on the weight of the two outer layers.

[0116] Aspect 5: The dry drug delivery system of any one of Aspects Ito 4, wherein the system is biodegradable.

[0117] Aspect 6: The dry drug delivery system of any one of Aspects 1 to 5, wherein the bioactive substance is a water-soluble drug selected from a small molecule of molecular weight < 1000 Dal, a large molecule of molecular weight > 1000 Dal, a protein of molecular weight between 10,000 Dal and 250,000 Dal, an amino acid, a peptide of 2-100 amino acid residues, or a nucleic acid.

[0118] Aspect 7: The dry drug delivery system of Aspect 5, wherein the nucleic acid is a plasmid DNA, a linear DNA, an oligonucleotide, an mRNA, or an siRNA.

[0119] Aspect 8: The dry drug delivery system of Aspect 1, wherein the bioactive substance is water-soluble and selected from the group consisting of an immunoglobulin, an antigen-binding domain, an anti-bacterial, an anti-viral, a steroid, an anti-glaucoma agent, an NSAID, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an anti-allergic, a lubricant, an antifungal, an anti-inflammatory, a protein, a peptide, a growth factor, an enzyme, a vitamin, a hormone, a polysaccharide, a peptide-nucleic acid, and a nucleic acid, or a mixture thereof.

[0120] Aspect 9: The dry drug delivery system of Aspect 1, wherein the bioactive substance is hydrophobic and selected from an anti-microbial, a steroid, an NSAID, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an antiallergic, a lubricant, an antifungal, an anti-inflammatory, and an anti-glaucoma agent.

[0121] Aspect 10: The dry drug delivery system of any one of Aspects 1 to 9, further comprising water-soluble adhesive fibers deposited onto the system, the water- soluble adhesive fibers selected from polyacrylic acid, PEG-multi-armed amine, polyethyleneimine, polyamidoamine, polylysine, polyarginine, chitosan gluconate, and derivatives and blends thereof.

[0122] Aspect 11: The dry drug delivery system of any one of Aspects 1 to 10, wherein the one or more excipients in the core of the core- shell fibers interact with thebioactive substance as a water substitute, complex with the bioactive substance, or form an interfacial layer between polymer-containing organic and water phases during fabrication of the system by electrospinning.

[0123] Aspect 12: The dry drug delivery system of Aspect 11, wherein the one or more excipients are selected from the group consisting of hydroxypropyl [3- cyclodextrin, gamma cyclodextrins, sulfobutyl P-cyclodextrins, PEG40-stearate, trehalose, sorbitol, sucrose, mannitol, poloxamer 407, poloxamer 188, polysorbate 80, polysorbate 20, dextran, polyvinyl alcohol, polyethylene oxides, polyvinyl pyrrolidone, hydroxypropyl methyl cellulose, human serum albumin, and soluble collagen, the one or more excipients being present in the dry drug delivery system in a concentration of 1-75% w / w.

[0124] Aspect 13: The dry drug delivery system of Aspect 1, wherein the system has been sterilized by cold e-beam radiation at a dose range between 6-25 kGy.

[0125] Aspect 14: The dry drug delivery system of Aspect 1, wherein the system is round, oval, ere scent- shaped, elliptical, annular, square, rectangular, or tubular.

[0126] Aspect 15: The dry drug delivery system of Aspect 1, wherein the hydrophobic elastomeric polymer forming the outer shell of the core- shell fibers is selected from the group consisting of ethylene vinyl acetate (EVA), polydioxanone, polysiloxane, P(TMC), and PCL, and the P(TMC)-co-PCL is present in the two outer layers at a concentration of 50-95% w / w based on the weight of the two outer layers.

[0127] Aspect 16: A method for delivering a bioactive substance to a soft tissue of a subject, the method comprising: obtaining a dry drug delivery system of Aspect 1, and applying the system to a surface of a soft tissue of a subject before or after adding a hydrating fluid to the soft tissue, the hydrating fluid comprising an isotonic solution that contains one or more of sodium chloride, a phosphate, a citrate, albumin, a magnesium salt, a calcium salt, borate, boric acid, a balanced salt solution, multi-branched PEG-amine, PVA, multi-branched polylysine, multi-branched polyarginine, guar gum, gellan, sodium alginate, xanthan gum, carboxymethyl cellulose, the hydrating fluid having a pH of 6-7.8 and an osmolality of 270-340 mOsm / kg.

[0128] Aspect 17: A water-dissolvable dry device, the device comprising a physical mixture of a first fiber and a second fiber different from the first fiber, the first fiber containing sodium hyaluronate, polyethylene glycol (PEG), and polysorbate 80 or polysorbate 20, and, optionally, castor oil, a buffer salt, an amino acid, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinyl pyrrolidone, dextran 70, albumin, PEG-multi- arm amine, collagen type 1, glycerol, or PEG40-stearate, and the second fiber containing polyvinyl alcohol (PVA) and sodium alginate and, optionally, sodium borate, dextran 70, PEG, PEG-multi-arm-succinimidyl ester, polypropylene glycol, propylene glycol, collagen type I, a bioactive substance, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or beta 1>6 glucan, wherein the first fiber and the second fiber are configured to react with each other and with a biological tissue upon hydration, and the device transform into a hydrogel upon contact with the biological tissue and subsequent hydration and completely dissolves over a period of 0 h to 14 days after contacting the biological tissue.

[0129] Aspect 18: The water-dissolvable dry device of Aspect 17, wherein the PVA has a molecular weight of 88,000-130,000 Dal with degrees of hydrolysis between 88-100%, the sodium hyaluronate has a molecular weight between 50,000- 5,000,000 Dal and the sodium alginate has a molecular weight between 32,000- 400,000 Dal, the first fiber including 80-90% by weight PEG, 0.5-5% by weight sodium hyaluronate, 0.1-0.3% by weight sodium chloride, 0.001-2% by weight polysorbate 20 or polysorbate 80, and the second fiber including 80-95% by weight PVA.

[0130] Aspect 19: The water-dissolvable dry device of Aspect 17 or 18, further comprising a bioactive substance selected from a small molecule of molecular weight < 1000 Dal, a large molecule of molecular weight > 1000 Dal, a protein of molecular weight between 10,000 Dal and 250,000 Dal, an amino acid, a peptide of 2-100 amino acid residues, or a nucleic acid.

[0131] Aspect 20: The water-dissolvable dry device of Aspect 19, wherein the bioactive substance is water-soluble, minimally water-soluble, or water-insoluble.

[0132] Aspect 21: The water-dissolvable dry device of Aspect 17, wherein, upon placement on a biological tissue, the device dissolves by simple solubilization in aqueous fluids or by biodegradation.

[0133] Aspect 22: The water-dissolvable dry device of any one of Aspects 17 to 21, wherein the device is round, oval, crescent- shaped, elliptical, annular, square, rectangular, or tubular.

[0134] Aspect 23: The water-dissolvable dry device of Aspect 17, wherein the device has been sterilized by cold e-beam radiation at a dose range between 6-25 kGy.

[0135] Aspect 24: The water-dissolvable dry device of Aspect 19, wherein the first fiber or the second fiber includes the bioactive substance, the bioactive substance being selected from an immunoglobulin, an antigen-binding domain, an anti-bacterial, an anti-viral, a steroid, an anti-glaucoma agent, an NSAID, an UV blocking agent, a healing agent, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an anti-allergic, a lubricant, an analgesic, an antifungal, an anti-inflammatory, a protein, a peptide, a growth factor, an enzyme, a vitamin, a hormone, a polysaccharide, a peptide-nucleic acid, a nucleic acid, or a mixture thereof.

[0136] Aspect 25: The water-dissolvable dry device of Aspect 21, wherein the biological tissue is the ocular mucosa, vaginal mucosa, rectal mucosa, nasal mucosa, or oral mucosa.

[0137] Aspect 26: A method for delivering a bioactive substance to a soft tissue of a subject, the method comprising: obtaining a water-dissolvable dry device of Aspect 17, and applying the system to a surface of a soft tissue of a subject, wherein the device transform into a hydrogel upon contact with the surface of the soft tissue upon hydration.

[0138] Without further elaboration, it is believed that one skilled in the art can, based on the above description, utilize the present invention to its fullest extent. The following specific examples are, therefore, to be construed as merely illustrative, and not limitative of the remainder of the disclosure in any way whatsoever. All publications, including patent documents, cited herein are incorporated by reference in their entirety.EXAMPLESExample 1: Dissolvable Alginate / Polyvinyl alcohol Inserts

[0139] This is an example of a crosslinked dissolvable insert formed from two solutions having different compositions. Solutions 1 and 2 were electrospun as two different fibers.

[0140] Table 1: Polymer Solution #1

[0141] Table 2: Polymer Solution#2

[0142] Table 3: Composition of Dry Device Insert

[0143] In a dissolution study in phosphate buffered saline at pH 7.4 and 37°C, the insert crosslinked upon hydration (within 5 minutes of soaking in buffer) turning into a tissue-conforming hydrogel. The hydrogel dissolved in 21-36 hours. Compositionally, the insert was designed to have both the alginate (crosslinked by Ca2+) and PVA (crosslinked by borate) form a longer-lasting interpenetrating network. The dissolution time can be decreased by lowering the boric acid concentration, and the calcium chloride concentration. In the absence of calcium chloride and borate, the PVA / Alginate / PEG35K insert dissolved in less than 8 hours ex-vivo. In another example, a PEG35K only insert was electrospun, the bulk of the insert dissolved in 30 minutes.Example 2: Dissolvable Hyaluronic acid / Alginate / PVA Insert

[0144] In this example, a hydrogel-forming insert was prepared by the process of electrospinning, by co-spinning two fibers simultaneously from two different solutions.

[0145] The insert was prepared using sodium hyaluronate EP2.0 (Freda Bloomage, Inc.), polysorbate 20 / 80, PVA88 (130,000 MW) (Sigma-Millipore), sodium chloride, and polyethylene glycol (20,000-35,000 MW) (Sigma-Millipore).

[0146] Solution 1 contained polyethylene glycol 35,000 MW (48%), boric acid (0.052%), calcium chloride (0.049%), polysorbate 80 or polysorbate 20 (0.21%), sodium chloride (0.41%) and endotoxin-free water. Solution 1 had a pH of 5.5-6.0, conductivity of 1.154 mS / cm, surface tension 31.59 mN / m, viscosity 9264.4 cP.

[0147] Solution 2 contained PVA 88 (8.97%), sodium alginate LVP (0.5%), polysorbate 20 (0.20%), acetic acid (0.146%), sodium chloride (0.40%), sodiumhyaluronate (0.5%) and endotoxin-free water. Solution 2 had a pH of 3.9-4.2, conductivity of 5.23 (mS / cm), surface tension 35.12 mN / m and viscosity of 1918.8 cP.

[0148] The final dry insert had a composition (by weight) of 28% PEG35K, 0.03% boric acid, 0.03% calcium chloride, 1.51% polysorbate 20, 61.15% PVA88, 3.4% sodium alginate, 2.93% sodium chloride, and 3.7% sodium hyaluronate.

[0149] The insert formed a hydrogel in 1 minute and became entirely transparent in 30 minutes. The insert was present as a thin hydrogel at 4 hours and dissolved over 4-6 hours. The use of calcium chloride in Solution 1 caused light crosslinking of the sodium alginate in the insert and boric acid in Solution 1 caused the PVA in Solution 2 to crosslink. The modulation of calcium chloride and boric acid in the insert composition can lower the crosslink density of the insert and the lower the dissolution time consequently. A control insert formed with no calcium chloride and boric acid dissolved quickly in less than 3 hours.Example 3: Dissolvable Collagen / polyvinyl alcohol Based In-Situ Forming Mucoadhesive Inserts (using a hydrating solution)

[0150] In this example of a dissolvable insert, the insert included Type I collagen fibers, PVA, and polysorbate 20. Upon placement on tissue, the mesh insert can be crosslinked in-situ to form a tissue-adherent hydrogel by application of a borate-phosphate saline hydrating solution. This insert was designed to dissolve while on tissue. The hydrating solution components trigger the crosslinking reaction. Using an electrospinning process, inserts were prepared from a single solution that contained Type I Collagen (4-8%), PVA 130K, 88% hydrolyzed (50-90%), Polysorbate 20 (0.1-2%) (see Table 4 below).

[0151] Table 4. Solution Composition of SB-NANOM-X-11-34

[0152] In order to generate a solution of collagen in any composition a pH <3.5 must be generated using acetic acid. The pH of the solution used was 3.3. Conductivity: 1875 pS / cm; Viscosity: 571.6 cP. Post-electrospinning, the PVA / collagen mesh insert is crosslinked in situ on tissue by use of borate in a hydrating solution. The hydrating solution was at a pH of 7.4-7.5, generated by the use of a borate-phosphate buffer.

[0153] The concentration of the small molecule drug can be 0.1%-10% in the dry insert, with PVA content ranging from 50-95%, and collagen ranging from 3-8%. The polyvinyl alcohol MW can be varied from 23,000 Da to 200,000 Da, and degree of hydrolysis varied from 88-100%. A 100% hydrolyzed PVA has no vinyl acetate content, and has very low solubility in water. Polyvinyl Alcohol was purchased from Sigma-Aldrich; Type I Collagen was purchased from Biomatrix, Nutragen Cat# 5010 (6 mg / mL, in IN HCL, pH 3.5) and Fibricol, Cat# 5133 (10 mg / mL in 1 N HCL, pH 3.5).

[0154] The water-soluble compound used in these studies had water solubility >100 g / mL, with a MW<500 Da. The compound is ionizable at pH > pH 4.5. A hydrogel formulation was monolithically electrospun. The hydrogel was formulated to consist of collagen, which is biocompatible and integrates structural integrity, and PVA, which is used to facilitate electrospinning by adding viscosity to the solution

[0155] The dry mesh formulation is the final formulation of the mesh after the electrospinning process.

[0156] Table 5. Dry Mesh Formulation of SB-NANOM-X-11-34

[0157] Electrospinning: During the electrospinning process, the solution formulation fouled up the tip of the needle at the typical voltage range of the needle(<18 kV). As a result, only a few fibers formed, but, for the most part, the formulation was an electrospray. Once the voltage was raised to 19 kV, the Taylor Cone stabilized, with only slight gelling at the tip. However, fibers were produced.

[0158] Table 6. Electrospinning Parameters for SB-NANOM-X-11-34

[0159] Use of a Hydrating Solution: An isotonic borate-phosphate saline solution was used as the hydrating solution to crosslink the components of the hydrogel insert and tether the drug delivery system to the underlying tissue. The inserts were crosslinked with either 20 mM, 25 mM, or 40 mM borate buffer. The hydrating solution was applied first to the tissue, and an 8 mm insert applied to the tissue. Without use of the borate-phosphate buffer, the insert hydrates, but does not tether to the mucosal surface. Thus, a borate-phosphate hydrating solution is a necessary design element of the collagen / PVA-based fast dissolving system. Preparation of Hydration Solution: The wetting solution contained borate, while the phosphate component assists in pH stabilization of the solution. The borate ions crosslink with the PVA component in the insert formulation (Fig. 7). The crosslinking mechanism is imperative for bio-adhesion and structural integrity of the insert. Phosphate buffer was added to the wetting solution to stabilize the pH of the solution, as pH 6.8 is outside of the pKa range for borate buffer (pKa @ 25°C: 9.23, 12.74, and 13.80, effective pH range: 8.5-10.2). Sodium chloride was added to adjust the osmolarity of the solution, achieve final osmolality.

[0160] Table 7. 25 mM Borate-3 mM Phosphate Buffer Formulation

[0161] Table 8. 25 mM Borate-3.3 mM Phosphate Buffer Characteristics

[0162] Rapidly-Hydrating: The collagen / PVA-based insert is initially white and dry. After application of the insert, a 40 microliter drop of the hydrating fluid can be applied, which triggers the hydration process of the insert. In another embodiment of this concept, the hydrating fluid can be applied before or after the insert placement on tissue. The insert after placement, starts to hydrate, absorb water rapidly to transform into a hydrogel. The hydrogel appeared to "sink" into the tissue, molding into crevices. This example insert was 60 microns in thickness. The hydrogel was no longer visible on the tissue post overnight incubation in buffer. There were no noted differences in crosslinking rates between the strengths of the borate buffer, pH 7.8. The thickness of the dry device can be between 60 microns and 1 mm in thickness, with a preferable thickness between 60-1000 microns, with the most preferable thickness between 100-450 microns. For some applications, the thickness of the insert may be as high as 2000 microns, although these applications would be for non- ophthalmic indications. When the collagen / PVA containing insert was placed onto 30 pL of 25 mM borate / 3 mM phosphate hydrating solution in an in-vitro test, its initially smooth opaque, white insert (8 mm diameter) absorbed water to become a round hydrogel. The time it took for the insert to form a full hydrogel structure was recorded and is shown in Table 1. Each sample shrunk on initial contact with the hydrating solution, and turned into a hydrogel, which is the result of the crosslinking mechanism between PVA, glycerol, collagen and borate ion. The crosslinking reaction between PVA and borate ions is well known, but an interpenetrating network between PVA, glycerol and collagen forms hydrogels that can be formed in- situ, on a tissue surface. PVA-only inserts, when crosslinked with borate, collapse into "slime", with no integral structure. The texture of the PVA / collagen / glycerol insert is smooth, and can deliver the drug to the site and dissolves away into non-toxic components. Thus, the inclusion of collagen is a necessary design element. Time to "hydrogel" formation: The time to hydrogel formation is dependent on insert thickness. Sample KR-NANOM-X-10-61-02 was the thinnest with a thickness of 0.006 mm. Once the insert was placed on the hydrating solution drop, it only took 43 seconds to dissolvemost of the insert, leaving a small piece to crosslink and form a hydrogel. The rest of the samples with greater thickness took longer than 1 minute to form a hydrogel, and did not dissolve rapidly. See Table 9 below.

[0163] Table 9. Time taken for the mesh to completely turn into a hydrogel

[0164] Maintains shape: Each sample shrunk and floated on the surface of the borate buffer, and maintained a round shape. The inclusion of collagen is an important design feature in maintaining shape of insert. Collagen at pH > 7 is insoluble in water, forming a water-swelled hydrogel. Dissolution Studies: Dissolution of the insert was determined upon contact with a 25 mM borate / 3 mM phosphate saline solution (Fig. 8: Crosslinking Kinetics and Dissolution of Collagen-Containing Hydrogel Formulation). Then, the newly formed hydrogel was diluted with 20 mM phosphate buffer, which partially simulated the environment around the fornix conjunctiva. The samples were placed in a monitored incubator @ 37 °C at the following time points: 0.5h, Ih, 2h, 4h, 6h. The physical state of the solution and hydrogel insert was observed and reported, and the solution was assayed to quantify for drug. -100% of the drug was recovered from the hydrogel.

[0165] Dissolution of the hydrogel was assessed in 20 mM phosphate buffer (pH 7.4, 37°C) for up to 62 hours. The buffer simulated physiological conditions, which partially simulated the environment around the conjunctiva. The samples were placed in a monitored incubator @ 37°C for the following time points: 0.5h, Ih, 2h, 4h, 6h. Once the 20 mM phosphate buffer was added, the insert remained a hydrogel, and floated on the surface of the solution, having a lower density than water. The vials were placed in the incubator, and removed at their respective timepoints.

[0166] After 64 hours of incubation, there was a thin translucent film left in the 20 mM phosphate buffer, which was used to simulate the environment in ophthalmic tissues.

[0167] Fig. 8 shows the dissolution of collagen-containing hydrogel inserts. Dissolution kinetics may be faster in tear fluid which includes lipids, enzymes and mucin. It was also concluded that insert thickness correlates with 'time to hydrogel' times. See Table 9 above.Example 4: Ex-vivo and In-vivo Studies of a PVA / Collagen-containing Monolithic Electrospun Insert

[0168] The insert was formulated with the addition of acetic acid to make the solution more compatible with the collagen component, which was supplied in a 0.01M HC1 solution, and to add more conductivity. The solution was notably clearer and uniform. The model drug was water-soluble with a molecular weight < 500 Da.

[0169] Table 10. Solution Formulation of Lot# SB-NANOM-X-11-75

[0170] Table 11. Dry Mesh Formulation of Lot# SB-NANOM-X-11-75

[0171] Electrospinning Parameters: The insert was monolithically electrospun, which implies the mesh was spun with one syringe and needle which was comprised of a single solution. During the electrospinning process, it was noted that beaded fibers and droplets were expelled from the needle tip at the beginning of the process. The initial humidity (67%) was dropped to 35%. As a result of the change in humidity, droplets from the needle tip seized. The mesh appeared uniform with an encapsulation average (n=3, STD: 7%) of 103% recovery of drug compound. This demonstrates that the mesh was uniform in regards to the API loading.

[0172] Table 12. Electrospinning Parameters of SB-NANOM-X-11-75

[0173] Ex-vivo study:

[0174] Ex-vivo studies were conducted to determine the bio-adhesive and hydro-gelling properties of SB-NANOM-X-11-75 on a bovine eyelid. The sample was placed on the eye lid, and the borate wetting solution was applied (see Fig. 9: Hydrogel Insert Application) to activate the crosslinking mechanism between the PVA component of the insert, and the borate component in the wetting solution.

[0175] As shown in Fig. 10, a hydrogel formed approximately 20 seconds after the wetting solution was applied. Tweezers were used to nudge the tissue and the insert. The insert moved with the tissue, which showed sufficient bio-adhesion.

[0176] As also shown in Fig. 10, after 90 seconds of wetting, the insert still adhered to the tissue and maintained its round, translucent properties. After 40 minutes of application, the insert was a transparent hydrogel, which was still intact.Example 5: In Vivo Studies of SB-NANOM-X-11 -75

[0177] Upon application to the rabbit eye, the insert transformed into a hydrogel, and did not appear to cause discomfort to the animal for up to 6 hours, with no excessive blinking, rubbing, or discharge. However, the insert was not visible, meaning it may have intercalated or molded into the tissue. The insert was not visible at the 1-hour timepoint or the 6-hour time point. See Fig. 11: Application of Insert to Conjunctival Fornix; Fig. 12: Application of Wetting Solution; Fig. 13: Hydrogel Insert after Wetting; and Fig. 14: Insert Intercalated into Tissue after 1 hour in both Eyes).

[0178] Eot# SB-NANOM-X-11-75 appeared to lack structural integrity because it dissolved within the first hour of administration onto the conjunctiva fornix in the in vivo studies. It is possible that there was inefficient crosslinking due to theborate solution not being evenly or efficiently administered. Without the PVA-borate mechanism of action taking place evenly across the entire insert and the tissue, the hydrogel cannot reach its full structural integrity. Therefore, the hydrogel formulation was further optimized as well as the electrospinning process.Example 6: In-Vivo and Ex-Vivo Studies of Co-spun Dissolvable Insert

[0179] SB-NANOM-X-11-75 showed insufficient crosslinking and the insert dissolved within the first hour of in-vivo application. This could be the result of a lack of structural integrity of the matrix. Based on the results from the in-vivo study, the hydrogel insert was reformulated to extend the dissolution time of the insert.Example 7: Dissolvable In-Situ Crosslinkable PEG35K / PVA100 / PVA88 / Alginate Insert

[0180] This example is an in-situ forming hydrogel insert, formed when hydrated with tissue fluid. No hydrating fluid is needed for this insert. This insert post-placement hydrates slowly with the tissue fluid, triggering the interlocking, crosslinking mechanism which tethers the insert to the underlying mucosa; the insert is hydrogel-like, aiding in comfort and feel. The hydrogel-like insert is transparent, disappears into the tissue and is not visible to the naked eye. Depending upon the specific concentrations of alginate and polyvinyl alcohol, the rate of dissolution can be varied.

[0181] The insert device is formed by co-electrospinning of two solutions, one containing a drug at a concentration that is therapeutically relevant (1-35%), PEG35K (10-60%), calcium chloride (up to 0.5%), boric acid (0.01-5%) and the other containing 0.01-0.5% polysorbate 20, 0.40-0.50% sodium alginate and 7 - 25 % PVA1OO / PVA88. Glycerol may be included as a surface wetting agent, and other buffer salts such as phosphates, or acetates, or histidine or citrate may be utilized.

[0182] In a particular example, the nanofibrous insert was formed by cospinning a first solution containing 55% PEG35kDa, 0.14% boric acid, 0.05% calcium chloride, 0.01% polysorbate 20 in water for injection, or distilled water and a second solution containing PVA100 - 2.6%, PVA88 - 7.7%, Sodium alginate - 0.44%, polysorbate 20 - 0.1%, acetic Acid - 0.5%, drug - 10%. The viscosity of the first solution was 23,184 cP, Conductivity: 56.8 pS / cm, pH 6.2 and surface tension 45mN / m; the viscosity of the second solution was 2270 cP, conductivity 565 pS / cm, pH 4.3 and surface tension > 45 mN / m.

[0183] The dry insert composition was 15.8% PEG35KDa, 0.028% boric acid, 0.01% calcium chloride, 0.003% polysorbate 20, 10.3% PVA100, 31% PVA88, 1.78% sodium alginate, 0.42% polysorbate 20, glycerol 0.7% and 40% model compound.

[0184] In another example, the insert was formed by co- spinning a first solution containing 35-55% PEG35kDa (Merck lot# 57742792 924, cat# 8.18892.1000); 0.14% boric acid (Boric acid: Spectrum lot#: 2GE0219, cat# B0120, NF, EP, BP, JP Grade); 0.15% calcium chloride (Calcium Chloride (dihydrate): Millipore Sigma lot# A0300782 629, Cat#1.02382, ACS, EP grade), 0.1% polysorbate 20 (Tween 20; Croda, super refined. SR40800) in water for injection (Q.S.), or distilled water, and a second solution containing PVA100 (6%), PVA88 (3.2%), Sodium alginate (0.46%), polysorbate 20 (0.01%), acetic acid (0.5%), drug (6%). The viscosity of the first solution was 24,000 centipoise (CP), measured at 25°C and 40 shear rate (Anton Paar Rheometer MC92, Spindle: Part# 18163, diameter 25 mm, Angle: 1°); Conductivity: 279 pS / cm (Mettler Toledo model#), pH 4.63 and surface tension >50 mN / m, and the viscosity of the second solution was 24,000 cP, Conductivity: 225 pS / cm, pH 4.63 and surface tension >45 mN / m.

[0185] The dry insert composition was 19.9% PEG35KDa, 0.04% boric acid, 0.06% calcium chloride, 0.04% polysorbate 20, 28.2% PVA100, 15.2% PVA88, 2.13% sodium alginate, 0.05% polysorbate 20, 35% drug substance.

[0186] In general, the viscosity of solution 1 is in the range 15,000-30,000 cP at 25°C, conductivity in the range 50-350 pS / cm, pH 4-7. The viscosity of solution 2 is in the range 1100-2500 cP, conductivity in the range 500-1000 pS / cm, pH 4-8.

[0187] Tested in-vivo in rabbit eyes, the hydrogel insert was deemed comfortable, and slowly dissolved over time.Example 8: Co-Spun Dissolvable Insert

[0188] Since the solutions used to process SB-NANOM-X-11-91 were cospun, two phases (solutions) were required. Sodium alginate was added to the cospun formulation, and so was calcium chloride. However, after several trials, sodium alginate would not remain in solution with collagen, due to pH incompatibility.Sodium alginate (pH 5-9 in water) and collagen (pH 4 in HCl / water) did not form a stable solution due to the pH needed for either molecule to remain stable within an aqueous environment. Sodium alginate is a carbohydrate that crosslinks via a reaction with calcium ions, and the hydrogel formation mechanism is due to the formation of PVA crosslinks via a reaction with borate ions. Collagen was removed because it served no purpose for promoting crosslinking or bioadhesion. This insert prototype did not require wetting.

[0189] Table 13. Solution #! SB-NANOM-X- 11-90-1

[0190] Table 14. Solution #! SB-NANOM-X- 11-90-2

[0191] Table 15. Dry Mesh SB-NANOM-X- 11-91

[0192] Electrospinning of both solutions could be accomplished. The viscosity of the PEG phase (Solution#!, lot# -90-1; see Table 13 above) should be lowered to improve flow and relieve pressure on the pump for future spins.Encapsulation analysis of the insert (average n=3, 108% drug recovered) demonstrates mesh uniformity.

[0193] Table 16. Electrospinning Parameters to Produce SB-NANOM-X-11-91

[0194] Table 17. ECU / Needle Positions

[0195] Table 18. SB-NANOM-X-11-91 Dry Insert Physical Characteristics

[0196] The thickness of the inserts in lot -91 was in the range of 130 to144 pm, with a density range of 0.216-0.254 mg / mm3. The radii of the inserts were 3 mm.Example 9: Ex-vivo studies on SB-NANOM-11-91

[0197] Freshly excised calf bovine conjunctival tissue was procured and used within one day of excision to assess insert lot -91. After placement on the wet conjunctival tissue, hydrogel formation was observed. After initial placement, the insert hydrated slowly, which triggered the interlocking, crosslinking mechanism which tethers the insert to the underlying mucosa (see Fig. 15). Just as with prior ex- vivo studies, clean tweezers were used to nudge the hydrogel, and it adhered to thetissue and maintained a round shape. After 24 hours, the insert was not visible to the naked eye (see Fig. 16).Example 10: Longer Lasting Formulation: Dissolvable In-Situ Crosslinkable PEG35K / PVA100 / PVA88 / Alginate Insert

[0198] Longer lasting hydrogels can be formulated by the inclusion of glycerol in the solution that contained polyvinyl alcohol (Solution #1). In prototype SB-NANOM-X-15-14, a combination of PVA 100 (fully hydrolyzed) and PVA 88 (partially hydrolyzed) was used to determine if the hydrolyzed state of the PVA would affect insert dissolution. Glycerol was also included in the formula as a surface wetting agent. The prototype was co-spun, just as SB-NANOM-X-11-91 had been, meaning that it did not require a wetting solution for the crosslinking mechanism to occur.

[0199] Tables 19 and 20 below show the two solutions that were co-spun using the same process used to fabricate 11-91. However, the sodium alginate and CaCh components were increased to enhance crosslinking in 15-14, which could decrease the rate of dissolution. Process control characterization is shown in Table 21, and the dry insert composition is shown in Table 22.

[0200] Table 19. Solution #! SB-NANOM-X-15-14-01

[0201] Table 20. Solution #2 SB-NANOM-X-15-14-02

[0202] Table 21. Process Control Characterization

[0203] Table 22. Dry Mesh Formulation

[0204] The electrospinning parameters are shown in Table 23 below. It was noted that the PEG phase viscosity could have been a little lower to establish an improved flow that would relieve the pressure on the pump, but overall, the plume narrow, and the process was stable.

[0205] Table 23. Electrospinning Parameters of SB-NANOM-NANOM-X15-14Example 11: In-Vivo Studies ofSB-NANOM-X-15-14

[0206] Structural integrity of the insert devices of lot 15-14 was assessed to determine if altering the concentrations of certain components in the formulation would affect hydrogel dissolution when placed in the fornix of a rabbit eye, namely concentrations of PVA (polyvinyl alcohol) and sodium alginate content. General observations were made on redness, comfort, and irritation at the specified timepoints of T=1 hour and 6 hours.

[0207] The results of in-vivo studies in New Zealand White rabbits are shown in Figs. 17 and 18. The insert demonstrated: (a) rapid hydration of the insert to a transparent hydrogel, (b) formation of an adherent device to the conjunctiva, and (c) absence of grooming or squinting by the animal after placement demonstrating that the hydrogel device was comfortable. At both the 1 hour and 6 hour time-points, the inserts were completely dissolved and no remnants were observed.

[0208] During the in-vivo study, Prototype #2 (15-14) was applied to each eye of 1 rabbit (see Fig. 19). The 6-hour time point was the endpoint for the study. After application (see Fig. 20: SB-NANOM-X-15-14 Insert Application), there was very slight ocular discharge in the nasal comer of the ocular surface, but it was considered minor. Following sedation (after 6 hours), the hydrogel was found and was slightly opaque, gummy, and circular (see Fig. 21: Postmortem Ocular Surface, T= 6 hours). The results suggest that 15-14 (which was formulated with two PVAs, the major component being fully hydrolyzed) did not reach full dissolution in-vivo for up to 6 hours.Example 12: Longer Lasting Formulation: The Effect of Glycerol Added in the Formulation of a Dissolvable In-Situ Crosslinkable PEG35K / PVA100 / PVA88 / Alginate Insert

[0209] The goals for mesh# SB-NANOM-X- 15-22 was to (i) increase PVA content so that the PVA phase increased in density, which may help make a more uniform final mesh, (ii) adjust the PVA blend to PVA1OO / PVA88 (25:75). A 65:35 blend, which consists of more fully hydrolyzed PVA, caused the insert to dissolve much too slowly. Lowering the amount of fully hydrolyzed PVA may increase dissolution times of the insert (up to 6 hours; the current target is 3 hours), and (iii) decrease CaCh and sodium alginate levels at the concentrations seen in 11-91 to have less crosslinking, allowing the insert to dissolve by the 3-hour mark.

[0210] Table 24. Solution #1 (PEG Phase) of SB-NANOM-X- 15-22

[0211] Table 25. Solution #2 (PVA Phase) of SB-NANOM-X- 15-22

[0212] Table 26. Dry Mesh Formulation of SB-NANOM-X-15-22

[0213] Stability studies were performed on samples (Mesh lot# SB-NANOM- X-15-22) that had been exposed to e-beam sterilization. This experiment evaluated the effect of 15 and 25 kGy E-beam exposure on dissolvable inserts containing the drug.

[0214] All samples were analyzed after 1 month in 5°C, 25°C / 60%RH (relative humidity), and 40°C / 75% RH stability conditions. Samples exposed to 15 kGy e-beam are adequately sterilized without destruction of insert properties.

[0215] Table 27. Sterilization by Ebeam

[0216] As shown above in Table 27, insert characteristics were maintained after 15 kGy ebeam irradiation. In another test, the insert device was sterilized by gamma irradiation under dry ice at 9 kGy. Insert characteristics were maintained.

[0217] The dissolvable insert concept can be utilized to deliver drugs, or healing ingredients to the surface of the eyes to treat disorders of the ocular surface. The disorders of the ocular surface may be disorders of the cornea, such as keratoconus, dry eye, blepharitis, Sjogren's syndrome, corneal abrasions and tears, corneal keratopathy, bacterial and fungal infections, allergic conjunctivitis, prevention of inflammation post-cataract surgery. Ocular disorders may include blepharoconjunctivitis, chalazion, conjunctivitis, contact lens problem, corneal abrasion, corneal dystrophy, corneal edema, corneal erosion, corneal ulcer, dacryocystitis, ectropion, endophthalmitis, entropion, episcleritis, eye tumor, foreign body, fungal keratitis eye infection, glaucoma, Graves' ophthalmology, hypotony, keratitis, migraine, mucormycosis, neuroretinitis, ophthalmoplegia, optic nerve problem, orbital cellulitis, photokeratitis, scleritis, sinusitis, stye, surgery, ocular bum, trauma, uveitis, UV damage, glaucoma, corneal vascularization. The dissolvable insert can also be utilized as a way to treat Sjogren's syndrome of mucosal tissues of other routes, such as the nose, rectum, mouth, and vagina. The dissolvable insert may be used with or without drugs, including small molecule, proteins, peptides, hormones, adenovirus-based drugs, RNAi, CRISPR-CAS9. The insert can be utilized to treat diseases of the brain, specifically for oncology applications.Example 13: Preparation of Dissolvable Inserts with Water-soluble and Insoluble (Oil) Ingredients

[0218] These inserts were designed to lubricate the ocular surface, generating an air-oil-water interface and prevent the evaporation of water.

[0219] Table 28. Composition and characteristics of solution IG-NANOM-DSW-02- 75A

[0220] Table 29. Composition and characteristics of solution IG-NANOM-DSW-02- 77AExample 14: Preparation of a Mutually-Crosslinkable Dissolvable Dry Insert Containing a 39-Amino Acid Peptide

[0221] The dissolvable inserts described in this example are comprised of 2 different types of fibers, one fiber containing multifunctional PEG8arm-NHS and the other containing PEG8arm-NH2. The insert is a dry device, which absorbs physiological water from the soft tissue surface. Crosslinking is triggered at pH>7, which mutually crosslinks the fibers and the tissue, generating a slow-dissolving insert. The solutions were prepared as per the compositions in Table 30 below. In this example, two solutions, each of them containing a multifunctional polyethylene oxide (either PEG8-arm- Amine, or PEG8arm-NHS) were prepared. (NHS = succinimidyl ester; GAS=glutaramide succinimidyl ester) Each solution contained Ethanol, THF, Polyvinyl Alcohol and Gellan. It is estimated that the pH of each solution was 6-7. The PVA solution (10% in water) and Gellan solution (1.5% in water) were mixed under stirring at room temperature (PVA / Gellan 80% / 20% w / w) for 20 minutes. To one portion of the PVA / Gellan solution was added the API (Exenatide), as part of the preparation of Solution 1. This mixture was vortexed for 5 min. PEG- Amine dissolved in Ethanol / THF was added to the solution above and vortexed for 5 min to complete the preparation of Solution 2. To another portion of the PVA / Gellan solution was added PEG-NHS dissolved in Ethanol / THF and Tyloxapol. The mixturewas vortexed for 5 min to form Solution 2. Ethanol and THF are added to increase the dielectric constant (conductivity) of the solutions. It is estimated that the pH of each solution was 6-7. The solutions were loaded to 3 mL syringes and connected to the syringe pump that would deliver the solutions at the same rate. The distance between the syringes and tips was approximately 5 cm. The electrospinning equipment manufacturer was Tong Li Tech Co, Ltd; the model was TL-01. The equipment was set up for electrospinning, by connecting the electrical cords to the syringe tips to provide high voltage electricity. A collecting drum or plate was placed in the chamber and grounded. The parameters were as follows: Voltage 9.5kV, RPM of drum 30, pump flow rate 0.3 mL / hour, distance of tip and receiving drum / plate=10 cm. The syringe pump was started, and the electric voltage for electrospinning turned on. The solutions were spun into a continuous insert. The insert was dried in a vacuum oven overnight at room temperature.

[0222] Table 30. Compositions of Electrospinning Solutions 1 and 2

[0223] Wetting with Hydration Buffer: The hydrating solutions below are for dry device inserts that have fibers that are annular, with the outer layer being hydrophilic and capable of mutual reaction and the inner layer being hydrophobic, and containing the drug. The purpose of the hydrating solution is to: (a) wet the tissue surface to establish maximal contact between the insert and the tissue, (b) trigger thecrosslinking reaction between the polymeric amine (NH2) and the polymeric succinimidyl ester (NHS) incorporated into the insert, and / or proteins on the tissue surface. The high pH of the hydrating solution triggers the crosslinking reaction. Examples of hydrating solutions were prepared in the following compositions, some examples of which are shown in the table below. The hydrating solutions contained formulations of sodium borate, sodium chloride and gellan. The force required to "peel" the insert from tissue after bonding (1-30 minutes), is the Peel Force (N).

[0224] Table 31. Compositions of Hydrating Solutions

[0225] Other hydrating solution variants can contain 0.1-0.5%% xanthan,0.1%-0.5% guar, or 0.1-0.5% hydroxypropyl guar, 0.1-0.5% balanced salts, 0.1-0.5% polycarbophil, 0.1-0.5% sodium hyaluronate.

[0226] Table 32. Insert Compositions for Hydration and Peel Adhesive Strength

[0227] Solutions 1 and 2, shown in Table 32 above, were co-spun onto a rotating collector of an electrospinning apparatus. Solution 1 (loaded onto a 3 mL syringe) and Solution 2 (loaded onto a 5 mL syringe) were co-electrospun under the following conditions: Voltage: 9kV; Flow Rate: 0.2 mL / hour for both syringes; Rotation of collector drum: 40 RPM. YL-NanoM-1-32 was a co-spun insert, with each of the reactive PEGs (see Table 32, Solutions 1 and 2) in a different solution so that they do not prematurely react. Both solutions were spun into monolithic fibers, with the fibers each containing one of the reactive PEG components. The peptide was previously shown to be unreactive with the PEG-amine, so it was included in Solution#!. In the presence of the hydrating liquid, an electrophilic-nucleophilic reaction between the PEGs is triggered to form a tissue-bound hydrogel.

[0228] Quantification of Mechanical Adhesive Strength of Insert to Tissue

[0229] Mechanical adhesiveness of the hydrated insert to conjunctival tissue was assessed using a Shimadzu Mechanical Tensile tester. The tissue was fresh bovine conjunctiva. The insert was YL-NanoM-1-32. The dry insert was applied to fresh bovine conjunctival tissue, then hydrated with 40-50 microliters of the hydrating solution. The mechanical tester was attached to one end of the insert. The hydrated insert / tissue was secured on a horizontal substrate. The force required to pull apart the insert from the tissue was recorded in N (stress) / mm (strain) (Peel Test).

[0230] As shown in Fig. 22, complete bonding occurs approximately 30 minutes after application of the insert on the conjunctival tissue, then wetted with the hydrating buffer (YL-NanoM-pep- 1-59-6; see Table 33 below). The peel force (N) at 30 minutes was 0.028N-0.05N. The dry insert (with no hydrating solution) and the insert hydrated immediately after placement on the tissue had a peel force of 0.008N. This allows removal of insert if applied incorrectly. The electrospun insert transformed into a clear hydrogel. A comparative test of peel strength of the various hydrating solutions listed in Table 33 was conducted with the Peel Test Model.

[0231] Table 33. Peel Test Force (N) of Insert from Fresh Bovine Conjunctiva, as a Function of Hydrating Solution

[0232] Hydrating solution YL-NanoM-pep- 1-59-6 contained both sodium borate and gellan, and demonstrated the highest peel force of the insert from the tissue (Fig. 22, Right). The hydrating experiment was performed with YL-NanoM-1-32 (see Table 32).

[0233] Kinetics of Peel Strength of Insert to Bovine Conjunctiva, using a gellan / borate, pH 7.8 buffer

[0234] The kinetics of the peel strength of the insert to bovine conjunctiva was measured. The insert was cut to a 1x2 cm size (YL-NanoM-pep-1-79-3; see composition in Table 34 below) and applied to fresh bovine conjunctival tissue. Fifty microliters of the hydration buffer was added to the insert and peel strength measured over time. The hydrating solution was 30 mM sodium borate buffer and 0.25% gellan. The results are shown in Table 35 below.

[0235] Table 34. Composition of YL-NanoM-pep-1-79-3 (Dried Placebo Insert)

[0236] Table 35. Kinetics of Bonding of Hydrophilic Sheath Insert to Bovine Conjunctiva

[0237] It was observed that longer bonding time of the insert to the tissue provided higher peel strength. It was also discovered that placement of the dry insert first, followed by the hydrating solution, provided superior bonding to when the hydrating solution was administered first on the tissue.

[0238] Kinetics of Peel Strength of Insert to Bovine Conjunctiva, using polymer / borate-based buffers

[0239] Electrospun insert YL-NanoM-pep- 1-79-3 (see Table 34) was enrolled in an ex-vivo bioadhesion test (with Shimadzu Instron Tester) with fresh bovine conjunctiva. 8 mm insert patches were placed on the tissue, and allowed to bond, with no applied pressure. After application, 40 microliters of hydrating buffer was administered on the patch to allow the patch to hydrate. All tests were performed with 5 minutes of bonding time.

[0240] Table 36. Hydrating Buffer Compositions

[0241] Table 37. Peel Strength of Electrospun Insert from Bovine Conjunctival Tissue Using Gellan / Borate, Sodium Hyaluronate / Borate, Polycarbophil / Borate, pH 7.8.

[0242] The data in Table 37 above shows that the inserts hydrated with Sodium Hyaluronate / Borate, Polycarbophil / Borate at pH 7.8 had higher peel strengths than the inserts hydrated with gellan / borate buffer at pH 7.8.

[0243] Other polymers that can be used in the hydrating solution include xanthan / borate, polyacrylate / borate, gellan / borate, guar gum / borate, PolyQuad- 1 / Borate, Polylysine / Borate, phosphatidylcholine / borate, Chitosan / borate, Polyarginine / Borate, PAMAM / Borate, PEI / Boratc, Trilysine / Borate, Tetralysine / B orate. Mannitol, sucrose, trehalose can be used as optional additives. Tyloxapol / borate, polysorbate 20 / 80, poloxamers, octaxynol, triton-XlOO can be used as optional surfactants to add, in the concentration range 0.01-1%, depending on allowable FDA limits and requisite concentrations needed to achieve requisite functionality. The concentration of the polymers can be varied from 0.05% to 0.5%. Concentrations of polymers <0.05% are ineffective, while concentrations above 0.5% tend to form viscous gel solutions that are not suitable for dispensation through eyedropper bottles. Sodium chloride is added to balance solution to tonicity between 280-320 mOsm / kg.

[0244] In another exemplary insert, the mutually reacting polymers PEG10K- 8Arm-NH2 and PEG10K-8Arm-NHS were combined in a single solution, with PVA. This is normally difficult to do, due to the mutually reactive components of the solution. However, it is possible to minimize reactivity by mixing the solutions cold, and maintaining a relatively low pH. This solution is adjusted to a pH 5-5.5 withhydrochloric acid (see Table 38). This prevents premature crosslinking of mutually reactive polymers.

[0245] Table 38. Composition of the Solutions of Combined PEGs Insert

[0246] Implication of pH

[0247] The composition of the combined solution can be in the pH range 4- 5.5. At pH<4, degradation of PEG10K-8arm-GAS (NHS) occurs with hydrolysis of the ester linkages. At pH > 5.5, premature reaction between the PEG components occurs, and consequently, the solution cannot be electrospun. In the pH range 4-5.5, hydrolysable linkages in PEG10K-GAS are stable, rendering this an ideal pH range.

[0248] Implication of Concentration of PEG-components: The composition of the combined solution can be increased or decreased, for 0.5% to 10% in the solutions to be electrospun. The higher the activated PEG concentrations, the higher the "anchoring" to tissue and the longer the time it takes for dissolution of the hydrogel, after application of the hydrating solution. In the dry device insert, the concentration of the activated PEGs is between 5-50%.

[0249] 200 microliters of ethanol were added to 1 gram of the solution to facilitate electrospinning fibers. Composition of the dry insert for YL-NanoM-pep-2- 21 was 84.85% PVA 67K, 7.58% PEG8arm-NH2, 10K, 7.58% PEG8arm-NHS. The conditions for spinning were: Needle tip size: 20; Flow Rate: 0.5 mL / hour; Voltage: 9.5 kV.

[0250] Reaction of Borate and Polyvinyl Alcohol:

[0251] The reaction of borate with polyvinyl alcohol (PVA) is known. But the reaction taking place in-situ on tissue allows interlocking of the polyvinyl alcohol - borate interpenetrating network with the tissue to anchor the insert onto the tissue surface. Thus, the presence of both PVA and the PEGs in the insert composition allows for two different mechanisms of tissue bonding to occur. Peel strength measurements were carried out on hydrated insert YL-NanoM-pep-2-21. The insert was flexible in the dry state with a tensile strength of 1.2-6 N / mm. The insert wasapproximately -200 nm in thickness. A 1x2 cm portion of the insert was crosslinked using a borate buffer hydrating solution. The hydration solution was 0.25% gellan / 30 mM borate, pH 7.8. The insert was hydrated right on fresh bovine conjunctiva. After a 30-minute crosslinking time, the insert was subjected to the peel test model, using the Mechanical tester. The peel strength for hydrated YL-NanoM-pep-21 was 1.4-3 N, based on multiple measurements.

[0252] This enhancement of adhesiveness of the insert to tissue is extraordinary and addresses the challenge of retention of topical sustained release drug delivery systems. Thus, retention of the insert to the tissue surface for the duration of release is a critical quality attribute to the success of the drug delivery system.

[0253] In another experiment, PEG8arm-NH2 and PEG8arm-NHS were tested at higher concentrations to 15% w / v in acidified water, to evaluate the adhesive strength to tissue. The adhesive strength to tissue was enhanced to >6N, due to the higher crosslink density of the network formed between PVA, the PEGs and the endogenous proteins on tissue.

[0254] In another experiment, the molecular weight of the PEGs was varied in the insert, to 20K 8arm PEGs. The adhesive strength of the hydrated insert to tissue was measured by the peel test and confirmed to be > IN, as the minimum criteria of good adhesion.

[0255] In another experiment, pH of the solution was lowered further to 4.5 and 5. The adhesive strength of the hydrated insert was measured by the peel test and determined to be greater >3N.

[0256] The composition of the hydrating solution was evaluated for further optimization of peel strength, by increasing the borate concentration to 50 mM, to achieve high adhesive peel strengths. At 50mM, the adhesive peel strength was >3N.

[0257] Integrity of an Encapsulated Peptide in Insert:

[0258] In this experiment, it was determined if a peptide could remain intact during the electrospinning process. 10-20 mg of the drug-loaded insert was sonicated in 1 mL of an aqueous buffer, to dissolve all contents. Ethanol was added to precipitate the polymers, and the resulting slurry was centrifuged. The supernatantwas removed for HPLC analysis (RP C18 column; see HPLC method in Table 39 below) and analyzed for peptide content.

[0259] Table 39. HPLC Method for peptide

[0260] Measurement of encapsulated Peptide in Insert : 10-20 mg of the drugcontaining insert product was sonicated in 1 mL of an aqueous buffer, to dissolve all contents. Drug content was determined by HPLC analysis (RP C18 column) and analyzed for peptide content. Thus, peptide contained in the insert was 1.6-2.8%, or 16-28 pg / mg.

[0261] Table 40. Encapsulation of Peptide (API) in Insert

[0262] The composition of YL-NanoM-1-32 and -34 were as given inTable 32 above. Solutions 1 and 2 were simultaneously electro-spun to form the drug delivery system on the same substrate.

[0263] Peptide Integrity: The peptide extracted from the insert was also analyzed for degradation products, characterized as % oxidization. The peptide maintained its integrity throughout the encapsulation process, (see Fig. 23).

[0264] In-Vivo Evaluation of Dissolvable Hydrated Insert

[0265] The in-vivo behavior of the hydrated peptide-containing insert was evaluated. Fig. 24 shows a hydrated insert adhered to rabbit ocular tissue. The tested insert was identical in composition to YL-NanoM-pep-02-13 (see Table 41 below) except for the absence of any protein. The insert appeared translucent, was highly hydrated, and was well-adhered to the ocular mucosa. There was no redness, or chemosis present at T=0, or at the end of 5 days. Percent hydration of the insert was approximately 65%-75%.

[0266] In-vivo Irritation Assessment of Peptide-Containing Dissolvable Inserts

[0267] In vivo studies with peptide-containing inserts were conducted in 6 normotensive New England White rabbits. Sterile inserts were inserted (placebo inserts) into the conjunctival cul-de-sac of each eye of the rabbits. Baseline redness of each eye was assessed before and after insert placement. The rabbits were assessed twice daily for 5 days. General irritation assessments were made. There was no redness or erythema on any of the treated eyes after 5 days. The inserts were thus deemed biocompatible.

[0268] Irritation Testing of Dissolvable Inserts by EpiOcular Surrogate Model

[0269] Inserts with a composition similar to YL-NanoM-pep-1-32 (seeTable 32 above) except for the absence of any protein were tested for irritation using the Epiocular Corneal Construct Model (Cyptorex, LLC), and compared to benzalkonium chloride as the control. These encased hydrated placebo inserts were classified as non-irritant, whereas the benzalkonium chloride solution was deemed as highly irritating. The encasement on the inserts was a PEG-based hydrogel, while the inner core matrix was PLGA-based.

[0270] In-Vivo Retention Studies of Dissolvable Inserts

[0271] Placebo inserts 8-armed activated PEGs (PEG8K-NH2 + PEG8K- NHS) at concentrations 10% in slightly acidified water (pH 4.5) were tested in an in- vivo retention study in New Zealand White rabbits, n=6 for 14 days. The optimized borate buffer, pH 7.8 was utilized for hydrating the inserts on tissue. The hydrated inserts were retained for 7-10 days (the study ended at 14 days).

[0272] In another equivalent study, highly 8-armed functionalized activated PEGs (PEG10K8arm-NH2+PEG10K-8arm-NHS) with IgG-containing inserts inrabbits (n=3, or 6 eyes) remained adhered to the conjunctival surface for 21 days. In another study, 16-armed functionalized activated PEGs inserts could be retained on mucosal conjunctival tissue for greater than 2 weeks.

[0273] In-Vitro Release (%) of Peptide from Dissolvable Inserts

[0274] 20-40 mg samples were placed in 1 mL Float-a-lyzer tubes, suspended in 40 g of buffer and rotated at 37°C. Time point samples were taken over the duration of study and analyzed by HPLC. Fig. 25 demonstrates release of exenatide, a highly water-soluble 39-amino acid peptide from the electrospun inserts of varied composition. Insert YE-NanoM-pep-1-64 (having the same composition as YE- NanoM-pep-02-13 except with exenatide instead of BSA and IgG) demonstrated a much slower rate than control peptide diffusing from the dialysis tube, compared to peptide in an aqueous solution (buffer at pH 7.4).

[0275] The encapsulation of other peptides has been accomplished in the electrospun insert. The insert was fabricated using an electrospinning process, and encapsulated in a blend of PEGA (RG504+poly(TMC-LL), 50:50). The solvent was methyl acetate, and the polymer concentration was 20%. The peptide was Thymosin Beta-4 a protein that in humans is encoded by the TMSB4X gene. The human protein consists of 43 amino acids (SDKPDMAEIEKFDKSKLKKTETQEKNPLPSKETIEQEKQAGES (SEQ ID NO: 1) and has a molecular weight of 4921 g / mol. Post fabrication, the insert was dried in vacuum to remove residual solvents. Encapsulation (mg / G): 20 mg of the drug-loaded insert product was dissolved in 1 mL of a dissolution solvent, to dissolve all contents. 1 mL of an aqueous solution was added to precipitate the polymers. ImL of the slurry was centrifuged and the supernatant was removed for HPLC analysis (RP C18 column) and analyzed for peptide content. Peptide encapsulation is expressed in pg peptide / mg product. Peptide Integrity: The drug product was dissolved in an organic solvent and analyzed for degradation products characterized as % oxidation.

[0276] In-Vitro Release and 1 hour burst (%): (20-40 mg) samples were placed in 1 mL Float-a-lyzer tubes, suspended in 40 g of buffer and rotated at 37°C. Time point samples were taken over the duration of study and analyzed by HPLC. As shown in Fig. 26, varied release profiles were shown by three different prototypes. As demonstrated, a small water-soluble peptide can be released in a sustained fashionfrom the insert. Other peptides both hydrophobic and hydrophilic can be incorporated into the insert, including melatonin, estrogen peptides, collagen peptides, Epitalon, Thymosin B4 and Thymosin Beta4 fragment.

[0277] Insert Irritation Testing of Crosslinked Dissolvable Inserts: In vivo studies with insert YL-NanoM-pep-02-13 were conducted in normotensive New England White rabbits (Toxikon Corp, Bedford, MA). Sterile inserts were inserted into the conjunctival cul de sac of each eye of the rabbits. Baseline redness of each eye was assessed before and after insert placement. The rabbits were assessed for 2-3 hours. No redness, animal discomfort or irritation were noted in normotensive rabbit eyes for placebo inserts of maximum thickness, for the duration of the study.

[0278] Encapsulation of IgG and BSA in PEG-GAS / PEG-NHS Crosslinked Insert:

[0279] Bovine IgG (Sigma Aldrich) a protein of MW -120,000 g / mole, was encapsulated in a dissolvable insert. Analysis of the IgG demonstrated that the protein had been encapsulated intact, and that the IgG was released intact over time.

[0280] In another example, inserts were produced having both IgG (MW 120,000 g / mole) and Bovine Serum Albumin (BSA; MW 62,000 g / mole). The composition is shown in Table 41 below.

[0281] Table 41. Composition of Insert Containing Two Different Proteins

[0282] The in vitro release results are shown in Table 42 below. Both IgG and BSA were released in a sustained manner.

[0283] Table 42. In-vitro Release of Two Bioactive Proteins from DissolvablePEGGAS / PEGNHS InsertExample 15: Incorporation of Intact Proteins at High Concentrations into Inserts

[0284] These experiments demonstrate that certain mixtures of excipients stabilize proteins at very high concentrations, >100 mg / g. Proteins are labile molecules and are prone to aggregation at very high concentrations. Thus, its microenvironment is critical in its stabilization strategies. Proteins are also susceptible to aggregation and eventual degradation as a result of contact with organic solvents, especially in the presence of both an aqueous medium and organic solvent.

[0285] To achieve high protein loading in the electrospun insert and being able to sustain release, protein solution is loaded as the "core" aqueous solution, and a polymeric polymer is loaded as the "sheath" organic solution, in a core-shell electrospinning set-up. After electrospinning, the protein is incorporated in the core channel, surrounded by a hydrophobic sheath later. The compatibility of different solvents and polymers used to make the electrospun mesh insert prototypes with IgGprotein was evaluated. In the experiments below, excipients that stabilize proteins at high concentrations were identified.

[0286] A 100 mg / g IgG in 50 mM phosphate buffer pH 7.4; 10 mg / g PVP, PVA, and PEG in 50 mM phosphate buffer pH 7.4 were prepared. The organic / water / IgG solutions were prepared by the mixing of IgG (10 mg / g in buffer) 1:1 with 40% v / v ethanol, IPA, DMSO, butanone solutions. T=0 samples were analyzed by size exclusion chromatography, using UV-vis detection at 220 nm. The remainder of the formulations were placed in the 40°C incubator and sampled on day 4 by repeating previous step (T=4 days). Table 43 summarizes the compatibility solutions.

[0287] Table 43. Compatibility Studies

[0288] As shown in Table 43, PEGs (polyethylene glycols), PVA (polyvinyl alcohol), PVP (polyvinylpyrrolidone), PAA (polyacrylic acid), borate buffer, phosphate buffer, TritonX-100, ethyl acetate, DMSO, poly (LLA-TMC) were compatible with IgG. After lyophilization, and reconstitution with water to 100-130 mg / g, these excipients assisted in the stabilization of both BSA and IgG, enabling the formulation of proteins in high concentrations. In another study, 0.3-1 mg / g of IgG was incubated with TFA (trifluoroacetic acid) and polyacrylic acid. Analysis by SEC HPLC demonstrated -100% recovery.

[0289] In another study, 100 mg / g of BSA was incubated with 8.5 mg / g and 18 mg / g Arginine. -100% recovery was noted with the solutions that contained higher Arginine content. Other excipients HPpCD (80-200 mg / g) and PEG2KDSPE (1-10 mg / g), when incubated with BSA (100 mg / g), demonstrated 100% recovery of protein.

[0290] In another study, IgG between 160-240 mg / g yielded -100% recovery with 2-10 mg / g HPpCD, 18 mg / g Arginine and 160 mg / g of HPpCD in phosphate buffer solution. IgG (250 mg / g in phosphate buffer) was incubated in separate solutions with 20-40 mg / g HPpCD in buffer, 40 mg / g HPpCD and ethanol in buffer, or 40 mg / g mannitol in buffer was recoverable in the range 98-100% post- incubation overnight. In another study, the effect of Dextran, PEG40 Stearate, Tween 20 and Sorbitol was investigated (see Table 44 below). These excipient combinations stabilized high concentrations of IgG (100 mg / g) with percent recoveries between 98- 102%.

[0291] Table 44. Excipient Combinations of Dextran, PEG40 Stearate, Tween 20

[0292] Polymer-in-organic solvent combinations such as n-methyl pyrrolidone, Span 40, THF, anisole, dichloromethane, methyl acetate, ethyl acetate, butyl acetate incubated with IgG-in-buffer solutions do not degrade the protein, allowing very high concentrations of protein (100-150 mg / g) to be incorporated intothe electrospun insert. IgG and albumin dissolved in phosphate buffer (pH 6-7.5) at concentrations 50-150 mg / g and containing polysorbate 20 (0.1-1%), polysorbate 60 (0.1-1%), polysorbate 80 (0.1-1%), polyvinyl alcohol (0.5-100 mg / g), polyvinyl pyrrolidone (, polyethylene glycol 3350 (0.1-100 mg / g), soluble collagen (0.1-20 mg / g), lecithin (0.01-0.1 mg / g), PEG40 stearate (0.5-1%), sorbitol (1-20%), glucose (0.1-10%), sucrose (0.1-10%), trehalose (0.1-10%), hydroxypropyl cyclodextrins (10- 200 mg / g), sulfobutyl beta cyclodextrins (10-200 mg / g), hydroxypropyl methyl cellulose (10-20 mg / g), hydroxypropyl cellulose (10-20 mg / g), hydroxyethyl cellulose, tamarind seed polysaccharide (5-200 mg / g), beta glucan (5-200 mg / g), xanthan (5-20 mg / g), hydroxypropyl guar (0.1-10 mg / g), alginate (0.5-10 mg / g), hyaluronic acid (0.1-2%) are stable long-term.

[0293] Hydrophobic polymers that are compatible with IgG and can be spun into nanofibrous inserts incorporating the protein are poly (trimethylene carbonate) (1-20%), poly(caprolactone) (1-20%), poly (lactic-co-glycolic acid) (1-25%), poly (sebacic anhydride) (1-20%), polyanhydrides (1-25%), poly-orthoesters (1-25%), polyester-amides (10-25%), poly(silicone)s (1-30%), polyurethanes, polyesterurethanes, polyanhydride-urethanes, ethylene vinyl acetate polymers with 20-50% vinyl acetate content and blends thereof.

[0294] The stability of protein in alkaline conditions (pH 7-8) is maintained by decreasing the number of negatively charged amino acids (Asp, Glu) and increasing the number of neutral hydrophilic amino acids (His, Asn, Gin, and Arg). Amino acids (arginine, glycine, proline, glutamine, histidine, cysteine) in the range 0.1-10% as low molecular weight excipients such as sucrose and sorbitol has been shown to enhance stability of bovine serum albumin and IgG. Polyamino acids (polylysine, polyglutamate) at concentrations 0.1-5% in aqueous buffers (pH 6-7.5) can stabilize proteins (IgG, growth factors)

[0295] Other proteins that are stable in the combinations above are nerve growth factor, fibroblast growth factor, epidermal growth factor, recombinant antibodies, therapeutic antibodies, antibody-like scaffold proteins, glycoengineered proteins and immunoglobulins. Among therapeutic proteins, antibody-based drugs, Fc fusion proteins, anticoagulants, blood factors, bone morphogenetic proteins, engineered protein scaffolds, enzymes, growth factors, hormones, interferons,interleukins, thrombin, fibrin, fibrinogen, clotting proteins and thrombolytics can be dissolved at high concentrations and stabilized by these combinations of excipients in the amounts stated.

[0296] Other proteins that may be incorporated into the insert are interleukin receptor antagonists, immunoglobulins, VEGF-inhibitors, aldose reductase inhibitors, antihypertensives, antioxidants, growth factor agonists and antagonists, vitrectomy agents, adenosine receptor antagonists, adenosine deaminase inhibitors, glycosylation antagonists, anti-aging peptides, topoisomerase inhibitors, anti-metabolites, alkylating agents, oncogene activation inhibitors, telomerase inhibitors, antibodies or portions thereof, fusion proteins, tyrosine kinase inhibitors, ribonucleotide reductase inhibitors, cytotoxins, IL2 therapeutics, neurotensin antagonists, peripheral sigma ligands, endothelin ETA / receptor antagonists, antihyperglycemics, anti-glaucoma agents, antichromatin modifying enzymes, insulins, glucagon-like-peptides, immunosuppressive agents, tissue repair agents, essential fatty acids, and nucleic acids such as plasmid DNA, linear DNA, string DNA, peptide nucleic acids, antisense oligonucleotides, mRNA, siRNA, and RNAi.

[0297] The following peptides can be incorporated into the electrospun insert product: Exenatide, Cyclosporine, Bacitracin, Vancomycin, Daptomycin, Voclosporin, Glutathione, Oxytocin, Insulin, Glucagon, Bivalirudin, Buserelin, Corticotropin, Cosyntropin, Enfuvirtide, Eptifibatide, Glatiramer, Epitalon, Melatonin, pineal peptides, adiponectin peptides and derivatives thereof.

[0298] Other peptides that can be incorporated into the electrospun insert are defensin peptides, cell-penetrating peptides, anti-coagulating peptides, milk-derived peptides, antimicrobial peptides, collagen peptides, GLP-1 peptides, PPAR inhibiting peptides, adiponectin targeting peptides, WNT targeting peptides, EGFR targeting peptides, hormone peptides, oncology peptides, TAT peptides. Other peptides that have been incorporated into the delivery system has been pineal gland peptides including but not limited to Epitalon, Melatonin.Example 16: Preparation and Characterization of IgG-Containing Insert (Hydrophilic Core-Hydrophobic Shell)

[0299] The protein can be contained in within core-shell structures, without the top and bottom layers of a sandwich, OR with the top and bottom layers of a sandwich structure.

[0300] The goal was to generate defect-free inserts, with high protein loading and integrity. The IgG-containing insert was composed of two types of electrospun fibers, with one type containing a core-sheath fiber containing the protein. For the core-sheath fiber (much like a co-axial cable), the "sheath" is hydrophobic and polymeric and acts as the encapsulating barrier to slow down diffusion of the water- soluble protein (contained in the core), so that sustained release of the protein can be achieved. In the absence of the core-shell structure, protein release is fast (100% in 1 hour at 37°C) from monolithic fibers (monolithic=single fiber micro structure).

[0301] The core solution (aqueous) (prior to electrospinning) contains a hydrophilic polymer that can be selected from the following list: polyvinyl alcohol (1-30%), using molecular weights (23,000-150,000 g / mole), polyvinyl pyrrolidone (PVP) (1-20%), Polyethylene glycol 2000-10,000-distearoyl -phosphatidyl choline (PEG-DSPE; 1-50%), PEG40-Stearate (1-20%), Dextran (1-3%), Sorbitol (10-25%). The sheath polymer solution (organic phase) used were PDLLA-TMC, poly(TMC), or poly (TMC)-co-poly(caprolactone) dissolved at a ratio 50:50 with pol(caprolactone, intrinsic dissolved at 20% w / w, in methyl acetate / DMSO in the ratio 90:10, or PCL+PTMC-PCL (15-30%, in methyl acetate / DMSO). Span 40 was used in this formulation as the surfactant, to generate an organic-aqueous interface to minimize protein aggregation. Suitable molecules to generate an organic-aqueous interface are dipalmitoyl phosphatidyl choline (DPPC), distearoyl phosphatidylcholine (DSPE), phosphatidic acid, Tweens, Spans, ceramides, Pegylated lipid molecules.

[0302] The feed rate of the core solution during electrospinning is varied from 0.1-5 ml / h, whereas the feed rate of the sheath solution can be 0.1-10 ml / h. The relative humidity is varied between 20%RH to 40%RH, with a favorable and the temperature is varied between 22°C to 30°C. The distance between the collector and the spinneret is 10 and 18 cm, and the voltage differential between the spinneret and the collector is 10-18 kV.

[0303] Shown in Table 45 below are core-sheath compositions for a co-axial protein-containing product with protein in the hydrophilic core, and polymer in the hydrophobic sheath. Each of the solution compositions for the hydrophilic core and hydrophobic sheath are provided. The hydrophobic polymer content in the dry device is between 50-95% w / w. PTMC=polytrimethylene carbonate; PCL=polycaprolactone.

[0304] Table 45. Core-Sheath Compositions for IgG Drug Product (IgG inHydrophilic Core, Polymer as Encapsulating Membrane Sheath)

[0305] Table 46. Compositions and Process Conditions for High Concentration IgG-containing Coaxial Drug Product with BioadhesiveExample 17: Intact Protein is Encapsulated in Insert Product

[0306] As shown in Fig. 28, the standard (1 mg / g IgG standard), preencapsulation solution containing 11% IgG (core) and IgG extracted from the insert product showed similar levels of the protein monomer (-73-77%), and high molecular weight fraction (14-21%) and low molecular weight fractions (0.9-2.5%). The data demonstrate the encapsulation of high concentration of protein in the insert product.Example 18: Release of Intact Protein from Insert Drug Product

[0307] Fig. 29 graphically shows the release of intact protein expressed as % monomer and other species (expressed as HMW1, HMW2 and LMW). A comparison of the standard at 1 mg / g and the released protein at each time point to 192 hours (8 days) shows that protein released at each timepoint is intact, and matches the standard.

[0308] The data also demonstrate sustained release of a highly water- soluble protein (IgG) from an insert that is less than 250 microns in thickness. The IVRT data from insert shown in Fig. 29is a coaxially spun insert with (PCL and PCL-TMC (10:90) blended at a ratio of 50:50) as the sheath polymer dissolved in methyl acetate, with 0.1% span 40 as a and IgG in PEG40-stearate, 0.1% polysorbate 20 in 20 mM phosphate buffer, pH 7 at a concentration of 11% as the core solution. The data demonstrate that it is feasible to incorporate high concentrations of protein in the thin insert, without a resultant high aggregated fraction (HMW1,2). As shown in the figure, no low molecular weight degraded protein was observed (the peak at retention time -13 minutes was attributed to polysorbate 20). In contrast, meshes that were spun without polysorbate 20, or PEG40 stearate in the core protein solution demonstrated very low protein recovery (<20%), after encapsulation. In another example, an insert that contained sorbitol in the core solution with the protein, also demonstrated excellent protein recovery (>90%). Fig. 30also demonstrate the sustained release aspect of the protein released from the insert product.Example 19: Release of Intact Protein from Insert Drug Product

[0309] The data in Figs. 29 and 30 demonstrates the sustained release of a protein from an insert. Both inserts were co-axially spun using an electrospinning apparatus.

[0310] The protein was contained in the core solution that also contained hydroxypropyl beta-cyclodextrin, which likely complexed with the protein providingprotection from aggregation; the sheath polymer was polycaprolactone (PCL12) blended 50:50 with poly (trimethylene carbonate-polylactide) (90:10) co-dissolved in methyl acetate at a concentration of 20%, 0.1% span40. In contrast, aqueous protein solutions emulsified in an organic polymeric solution (comprised of the sheath solution from above), resulted in immediate burst of protein from the insert.

[0311] Fig. 30 shows sustained release from a poly(trimethylene carbonate)- co-poly(lactide) fibers. However, the inserts did not show dimensional stability, was highly stiff and thus not biocompatible on tissue. The release profiles on multiple inserts were not reproducible, more given to process nuances than release profiles obtained with the sandwich insert, which contained poly(TMC)-co- poly(caprolactone)(90: 10).

[0312] Fig. 31 shows high recoveries of protein from multiple batches using co-axial electrospinning of aqueous protein solution as the core, and an organic polymeric solution as the sheath. The data demonstrates the utility and reproducibility of protein-containing core-sheath insert products. In contrast, batches that did not contain PEG40-stearate, PEG2K-DSPE, 0.1% polysorbate 20, or sorbitol (0.1-11%), or 1-50% HPpCD had recoveries < 20% consistently. It demonstrated that the amphiphilic surfactants create a protective interface between the aqueous protein solution (core solution) and organic polymer solution (sheath), thus preventing denaturation events that lead to aggregation. Fig. 31 shows the relative recovery of different batches as a percent. Batches 12-995 A, 12-995B and 12-665C all had the excipients above, but at concentrations < 0.2%. This resulted in % recoveries that were comparatively lower.Example 20: Method, of Sterilization of Protein-Containing Insert

[0313] A critical requirement of a protein product for ocular administration, or administration within the body is the sterility requirement. Protein solutions used for injections and infusions are normally sterile-filtered unless high concentrations of proteins or excipients render the solutions too viscous to be sterile-filtered. The IgG- containing insert was produced with the components shown in Table 47 and was subjected to electron beam sterilization.

[0314] Table 47. Protein-Containing Insert Product for eBeam Sterilization Testing

[0315] Ebeam sterilization of protein-containing insert product was performed at 15 kGy, 25kGy, 40 kGy ebeam intensities (Ebeam Services, Inc.), while cooling the product on dry ice. After sterilization, the mesh inserts were analyzed for analysis of protein content, by size exclusion chromatography (SEC). The results are shown in Table 50 below.

[0316] Table 48. Protein Recovery as a Function of Electron Beam Sterilization

[0317] The insert shown in Table 47 contained a high concentration of IgG (9%) in the core solution with Tween 20, and sorbitol as protein-protective excipients. The sheath solution contained components that were also shown to be compatible with proteins. The data in Table 48 shows that ebeam exposure at 25 and 40 kGy was detrimental to the encapsulated protein (% protein recovery ~50%-68%, whereas exposure at 15 kGy demonstrated 96-100% protein recovery. Other sterilization schemes that were demonstrated to have retained functionality of protein was NO2 sterilization (Noxilizer, Inc.) (96% protein recovery), X-ray and Gamma Irradiation at doses between 1-15 kGy. Sterilization methods that did not retain functionality were gamma sterilization (<10% recovery) at doses > 15 kGy, and ethylene oxide (ETO). There were marked increases in low molecular weight fractions with high electron beam exposure (left, Fig. 32).Example 21: Selection of Dimensionally Stable Polymers to Construct the Drug- Containing Hydrophobic Fibers

[0318] Polymers comprising the hydrophobic fibers of the insert need to be able to maintain their size and shape after placement on the ocular tissue. An insert that does not exhibit dimensional stability, when placed onto the conjunctival tissue, can stretch and pull away from the tissue causing mucoadhesive failure and discomfort.

[0319] Experiment#!: Inserts were constructed of the following materials: (a) Polycaprolactone (PC12) blended 50:50 with poly (trimethylene carbonate): Polycaprolactone (90:10) (Sigma- Aldrich), (b) polylactide-co-glycolide, 5004 (Purac) (c) polydioxanone (PDO), (d) polycaprolactone, 80K (PCL 80K). Results: Upon hydration, the PLGA insert shrunk in volume and became stiffer in texture and feel. In contrast, PCL / PTMC blend (polycaprolactone 80K blended with 90:10 polytrimethylene carbonate: polycaprolactone) (Sigma Aldrich), PDO (polydioxanone, Sigma Aldrich) and PCL80K maintained dimensional stability, with zero percent of dimensional change upon hydration. The inserts upon hydration increased <20% in weight, indicating that these are low-swelling insert devices. Additionally, these inserts when extended could be pulled (strain) could be extended to > 100% its original size prior to cohesive failure demonstrating its elastomeric character. In contrast, PLGA inserts could not be extended to higher elongation(<10%), resulting in rapid cohesive fracture. Additionally, the PLGA inserts upon hydration, became highly stiff with sharp edges.

[0320] Analysis: For applications that require maximal contact with the underlying tissue surface, loss in dimensional stability results in delamination of the insert from the tissue surface. As shown in Fig. 33, PCL / PTMC blend, PDO and PCL80K maintained dimensional stability and low % swelling, making only those compositions suitable for ocular use, or for any tissue that requires a tissue-pliable drug delivery system for maximal comfort, such as nasal, rectal, vaginal or esophageal tissue. PLGA would not be suitable for any tissue applications that require flexible inserts / drug delivery systems.

[0321] Experiment#2: In another study (Table 49 and Fig. 34) PCL Blends (Polycaprolactone (PC 12) blended 50:50 with poly (trimethylene carbonate): Polycaprolactone (90:10)) were fabricated into mesh inserts of -30-50 microns in thickness and incubated in phosphate buffer, pH 7.4 at 37 °C initially for 24 hours, then for 72 and 14 days, respectively. The % dimensional change when hydrated was 0%; additionally, the % hydration was <20%.

[0322] Table 49. Time Course Study for PCL Blend Insert

[0323] In another study, PDLLA-TMC (50:50) (Sigma Aldrich) was fabricated into co-axial mesh inserts of -30-50 microns in thickness and incubated in phosphate buffer, pH 7.4 at 37 °C for 9 days. Measurements were taken at T=0 and T=9 days. Fig. 35 demonstrates that dimensional stability of inserts is an important product attribute. After 9 days, the PDLLA-TMC insert lost flexibility and shrunk many-fold (> 20% of its original size) into a sharp-edged, stiff inserts, inappropriate for soft tissue spaces. The co-axial mesh insert contained a protein (IgG) along- with protective excipients (phosphate pH 7, dextran, sorbitol, PEG40 stearate, cetyl alcohol) in the inner hydrophilic layer and PDLLA-TMC as the encapsulating hydrophobic layer.

[0324] Table 50. PDLLA-TMC Co-Axial Insert Dimensional Study

[0325] Polymers that are suitable for maintaining dimensional stability are poly (trimethylene carbonate): poly(caprolactone) copolymers at copolymer ratios of 90:10, 80:20, 70:30, blended in with polycaprolactone at molecular weights of 80K- 150K. Also suitable at polydioxanone / polytrimethylene carbonate / polycarbonate blends, at various ratios to modulate drug release. Another polymer that can be added in is poly ethylene vinyl acetate (EVA), to be added into a blend with polycaprolactone and poly (trimethylene carbonate), or poly(dioxanone) or used singly.

[0327] Experiment#3: EVA dimensional study

[0328] Table 51. Ethylene Vinyl Acetate-Based Insert Dimensional Study

[0329] Ethylene vinyl acetate inserts demonstrated high dimensional stability (% dimensional change < 10%, % hydration <100%). In addition, the percent elongational strain for ethylene vinyl acetate inserts >100%, indicating high elastomeric character. (% elongation: 20%, T=0 2.5 cm, T=after stretching 3.0 cm) Example 22: Development of a Bioadhesive Insert

[0330] In order to impart adhesion of the device to tissue, the drug-containing device can further contain water-soluble fibers of adhesive character that are codeposited onto the device, with the water-soluble fibers selected from the group polyacrylic acid, polyethyleneimine, polyamidoamine, polylysine, polyarginine, chitosan gluconate and derivatives and blends thereof. For use on mucosal tissue, the drug delivery system must be bioadhesive. In the example below, the bioadhesive molecule is polyacrylate, an anionic bioadhesive co-spun with a blend of poly(caprolactone) and a copolymer of poly (caprolactone-co-poly (trimethylene carbonate) (90:10). In other embodiments, the bioadhesive molecule may be polycationic, such as polyethyleneimine, polyamidoamine, polylysine, polyarginine, chitosan gluconate. A qualitative rating scale was developed to rate the bio-adhesion effectiveness of an insert when placed onto fresh bovine conjunctiva. The rating scale was designed by observing the characteristics of adhesion and appearance. Three different polyacrylic acid (PAA) molecular weights were rated for adhesion and appearance when placed onto conjunctiva tissue using a rating scale. 1.8kDa, 450kDa, and 3 million Dalton molecular weight polyacrylic acid (PAA) -containing inserts were tested individually and with other polymeric blends, to determine which formulation produces ideal adhesion to the bovine conjunctiva. The PAA blend of 1.8kDa and 450kDa at a ratio of 1:2 spun into the mesh produced ideal characteristics for the adherence of the insert to the bovine conjunctiva tissue. The insert did not slidearound on the tissue and peeled off with resistance, stretching the tissue, at timepoints 0, 24, and 48 hours.

[0331] Attributes of the Testing Method

[0332] A qualitative rating scale for the bio-adhesion effectiveness of the NanoM Insert when used on the conjunctiva region of bovine eyelid tissue was developed. The following criteria went in designing the adhesion rating scale: (a) adherence of insert to conjunctiva, (b) slide potential of the insert on the conjunctiva, (c) stretching of conjunctiva after removal of insert, (d) appearance of the insert on the conjunctiva, e) peel potential of the insert and (f) insert breakage during removal of the insert from the conjunctiva. Insert samples were prepared by using a sterile medical grade punch (Fig. 36). Each piece was carefully "punched" out from various portions of the electro-spun fiber mesh. Each Insert's dimensions were carefully measured and recorded into a notebook or spreadsheet. Initially, the inserts were cut using medical grade scissors into square or rectangular pieces. Fresh bovine conjunctiva tissue was received on the day of analysis. Once the tissue was received, it was placed inside the refrigerator on ice without any buffer until testing was ready to begin. Once the insert samples were prepared, the fresh conjunctiva tissue was removed from the refrigerator and placed onto a piece of wax paper. Each sample of tissue can be used for more than one test, there is an upper and lower eyelid, each containing a section of conjunctiva tissue.

[0333] Different molecular weight sizes and blends of PAA spun into PCL + PCL / PTMC blends produced different adhesion and appearance rating scores. Video recording began prior to placement of the insert onto the tissue. In the video, the lot number of the mesh, timepoint, and PAA content was verbalized for the record. A square or circular- shaped insert was carefully placed onto the freshly prepared conjunctiva tissue using forceps. Each tissue sample was kept on ice at room temperature until testing time, testing was done promptly after receiving the tissue to ensure fresh samples. During the recording of the video a verbal description of any changes to the appearance of the mesh was recorded, along with verbal analysis of adhesion performance. After a maximum of one minute, or three days if doing further timepoints, the rating for adhesion was scored by the resistance, or lack thereof, when removing the mesh from the tissue using forceps. When the mesh was difficult to peeloff of the conjunctiva, adhesion score increased (Fig. 37). The rating scale for adhesion and appearance can be referenced in Table 52 below. An adhesion score of 1 or more at tO qualified the mesh to be tested for adhesion at 24 hours (t24).

[0334] Table 52. Rating Scale

[0335] Description of Test Samples for Bioadhesion

[0336] Table 53 below summarizes bioadhesion testing of a 50:50 blend of polycaprolactone and poly (trimethylene carbonate: polycaprolactone 90:10) inserts, co-spun with polyacrylic acid as the bioadhesive. As shown in Table 55, polyacrylic acid was approximately 30% of the dry insert. Polyacrylic acid was varied between 1.8K and 450,000K and tested in various ratios of 1:1, 1:2, and 1:3 PAA (1.8K) and PAA (450,000K). Ratios of 1:2 and 1:3 yielded the highest bioadhesive scores. The insert that contained poly (acrylate) in a single molecular weight did not demonstrate strong adhesion, as the blends of different molecular weight PAA. This was an unexpected result, as all polyacrylates are known to be bioadhesive, agnostic of molecular weights. In contrast, non-poly aery lie acid containing inserts did not demonstrate appreciable adhesion to tissue.

[0337] Other polymers present in the dry insert that can impart bioadhesion are guar gum (1-30%), alginate (1-30%), polylysine (1-30%), polyarginine (1-30%), xanthan gum (1-30%), chitosan (1-10%), polyethyleneimine (1-10%). In another experiment, polyarginine was incorporated into the insert samples (1-10% polyarginine) comprised of a 50:50 blend of polycaprolactone and poly (trimethylene carbonate-co-polycaprolactone 90:10), formed by an electrospinning method. When hydrated, the insert demonstrated high adhesion (score=3). Likewise, Poly lysine (1- 10%) incorporated into insert samples of the above blend also produced strong adhesion scores of 3.

[0338] Table 53. Summary of Bioadhesion study of mesh containing 50:50 blend of PCL and PCL-P(TMC) (10:90)

[0339] All of the insert products described above were biodegradable, albeit at a rate that may be slower than the rate of drug release. Thus, these inserts may need to be removed, in order to receive the next insert product. In another experiment, Ethylene vinyl acetate (EVA) was utilized as the encapsulating hydrophobic polymer and polyarginine was used as the adhesive polymer. The adhesion score was 2-3.

[0340] Examples of "SANDWICH" INSERTS

[0341] Sandwich Inserts: These are compositions of electro-spun fiber-based inserts focused on (a) encapsulation of concentrated levels of highly water-soluble and highly labile molecules, (b) methods to release said molecules in sustained manner, (c) minimization of irritation at the local tissue and (d) maximization of retention at the local site of delivery.

[0342] The sandwich inserts have a hydrophobic barrier to drug release, in the form of a hydrophobic "shell" layer encircling a hydrophilic "core" layer, which contains the drug. In order to accomplish a linear rate of release, a "sandwich" insert has been described, with the drug-contained fibers (core- shell, or monolithic) sandwiched by two membrane layers. The inserts described, have a hydrophobic barrier to drug release, in the form of a hydrophobic "shell" layer encircling a hydrophilic "core" layer, which contains the drug. In order to accomplish a linear rate of release, a "sandwich" insert has been described, with the drug-contained fibers (core-shell, or monolithic) sandwiched by two membrane layers.

[0343] Without the shell layers (top and bottom layers), the inserts demonstrate high % burst of protein. This implies the core-shell fibers with the protein contained in the core, do not adequately provide a barrier to provide a sustained release of the protein. While Fig. 30 (Sustained Release of IgG) did show sustain release, this result is difficult to achieve in a reproducible manner. Often, the % burst of protein is >70%, with 100% released within 1 day. To solve this issue, the sandwich inserts were designed and implemented, with the core-shell fibers (containing protein) encapsulated further by two outer layers. The top and bottom layers need to be composed of polymers that "flow"; poly(trimethylene carbonate) - based polymers are known to have flow properties to completely encapsulate the inner core-shell fibers. To enhance the elastomeric character of the poly(trimethylene carbonate)-based fibers (both core-shell, and encasing), ethyl vinyl acetate can be added in the formulation. In one example, addition of ethyl vinyl acetate to the formulation enhanced the elastomeric elongation to greater than 500%.

[0344] The selection of the polymer for the encasing layers (top and bottom layers) and the processing conditions post-fabrication are critically important in achieving the design attributes of the sandwich insert. Only polymers that have molecular flow properties at temperatures 28-36°C are useful in generation of the encasing layers. The benefit to using electrospinning and not solvent casting is that the latter would result in high concentration of organic solvent in the drug delivery device, which degrades the protein. The process of electrospinning evaporates off the volatile organic solvent prior to device fabrication. Post-fabrication of the electrospun insert, the polymers comprising the layers flow to create continuous top and bottom layers.Example 23: Insert Lot no: AB-NANOM-X-07-65

[0345] Core: Core is an aqueous based solution containing a buffer, sugars, surfactants (tween 20, PEG, poloxamer), salt and the protein. All the excipients help in lowering down the surface tension and protect the protein from degradation. The preparation of the core solution is explained step by step below:

[0346] Prepare Histidine buffer: To prepare Histidine buffer, use a sterile container and weigh the required amount of WFI. On an analytical balance, weigh L- histidine and L-histidine HC1 and add it to the container. Add a magnetic stir bar to the container and place it on a magnetic stir plate. Once the excipients are in solution, measure the pH. Target pH 6.0-6.4. Adjust if necessary. See Table 54.

[0347] Prepare excipient stock solution: To prepare the excipient stock, in a new sterile container, weigh the required WFI followed by weighing the histidine buffer. Add a magnetic stir bar to the container and place it on a magnetic stir plate. On an analytical balance, weigh out all the excipients except HpPCD / cyclodextrin. Add all the weighed excipients to the sterile container. Once everything is mixed properly, weigh out HpPCD / cyclodextrin and add it to the container. Eet the solution mix properly till it's clear. Filter the excipient stock solution using a sterile 0.22 pm PVDF syringe filter into a new sterile container. See Table 55.

[0348] Prepare final core solution: To prepare the core solution, use the excipient stock solution directly to dissolve the protein (lyophilized protein stored in glass vials). Add the target amount of excipient stock solution to the glass vial and swirl to ensure proper dissolution (don't vortex). Once the protein is dissolved,remove the solution from vial and transfer it to a different vial. Repeat this step for all the protein containing glass vials (the no. of vials depends on the target cone.). After all the protein is dissolved properly, transfer the solution to a 5 mL sterile centrifuge tube and store it at 5 °C. See Table 56 for the composition.

[0349] Table 54. Preparation of histidine buffer (SS-NANOM-X- 11-64)

[0350] Table 55. Excipient Stock (SS-NANOM-X- 11-64)

[0351] Table 56. Core Solution (SS-NANOM-X- 11-64)

[0352] Core characteristics: Visual: Clear or slightly turbid solution (depends on the protein concentration); pH: Target 6.0-6.4; Conductivity: <2.5 mS / cm; Surface tension: 30-34 mN / m

[0353] Core characterization: Surface tension is an important measurement parameter for the core solution, as it determines how smooth the coaxial electrospinning will be. The surface tension is measured by using the tensiometer (Biolin Scientific Attention Theta Lite). pH is essential to determine the stability of the protein, if not in range can result in protein degradation. A pH meter (Fisher brand Accumet AB315 or equivalent) is used for a precise measurement. pH paper strips can also be used depending on the sample size; Conductivity is another parameter that determines the success of the electrospinning process. Mettler Toledo AG Five Go or equivalent can be used; The core solution is quantified using the UPLC method. This determines the loading of the insert. Sheath solutions: For the sandwich insert, there are two sheath solution which are used: Coaxial sheath and Film sheath. The coaxial sheath is the polymer solution which encapsulates the core solution, also called the middle layer. The film sheath is used in the top and the bottom layer. Both these solutions together make the sandwich mesh / insert.

[0354] Method of preparation: Weigh out all the solids in separate weigh boats for both the solutions; After all the solids are weighed out, tare the glass vial for the first solution on an analytical balance. Use a 3 mL syringe and a 25 G needle to pull out 3 mL of DMSO from the DMSO aliquot vial. Weigh out the required amount of DMSO in the tared vial; Tare the vial again and add all the solids required for that solution. Tare the vial again. Use a 10 mL syringe with 25 G needle and pull out at least 10 mL of Methyl acetate from the sealed vial. Add Methyl acetate to the tare vial and record the weight. Repeat the above steps for the second solution (film sheath). Label the vial properly listing out all the solvents and polymers; Place both the vials on the rotisserie in an incubator set at 40-42 °C. Let the solutions mix overnight.

[0355] Table 57. Composition of Coaxial Sheath Middle Layer

[0356] Table 58. Composition of Outer Encasing Layer

[0357] Sheath characteristics

[0358] Visually, the sheath was a clear solution with no undissolved polymer. The viscosity was measured with an AntonPaar MCR92 Rheometer (cone and plate). A viscosity of 1500-4000 cP is needed for successful electrospinning.

[0359] Table 59. Inherent Viscosity of Primary Encasing Polymers

[0360] Table 60. Electrospinning parameters

[0361] For P2, the encasing layer and the core-sheath solutions are co-spun.

[0362] Encapsulation procedure. This process is used to analyze the encapsulated protein / drug inside the insert. As the top and the bottom layer consist of polymers, organic solvents are used to dissolve the polymer. For this method, 50:50 Methyl Acetate: Dichloromethane is used to perform multiple washes to dissolve the polymer. After the washing process, the protein pellet is dried to remove left over solvents. This step is followed by reconstituting the dry pellet using 50 mM phosphate buffer, 1.2% Sodium chloride, 0.1% Poloxamer. The insert is analyzed by using a UPLC method. Sampling the Mesh for Encapsulation: Use a sharp, sterile 10 mm dermal punch to remove 1 punch from the mesh from the center, avoiding the stabilization zones; Place the sample inside a tared 1.5 mL centrifuge tube, and record the weight of the sample (the sample weight should weigh ~10 mg). Separating the Polymeric Components from the Aqueous Components: Using a syringe and a needle, extract 0.4 mL of Dichloromethane and 0.4 mL Methyl Acetate form their respective vials and add it to the centrifuge tube; Vortex the centrifuge tube in five second intervals for 2-3 times to break the insert into small pieces; After the insert appears broken up, sonicate the tube in 5 seconds intervals for 1-1.5 min or till the insertdisperses almost completely; Once the insert is fully dissolved / dispersed, place the tube inside a temperature-controlled benchtop centrifuge for 30 minutes @ 13.2K rpm at 5 °C.

[0363] Protein Wash Procedure: Remove the sample from the centrifuge, and note the appearance. The suspended material will form a pellet, and the solvent phase should be clear; Use a 1000 pL pipet to remove ~0.6- 0.8 mL of the solvent, being careful not to disrupt the pellet or any small insoluble particles. Removing solvent removes the sheath components; Repeat step c to h three more times and add fresh solvents at each step. Skip step e, as sonication is only required in the initial dissolution; Drying the protein pellet: Leave the centrifuge tube cap open and tightly seal the centrifuge tube with parafilm. Use a sterile needle (26-30G) to pluck 3-4 holes into the parafilm. The holes should be big enough for the solvent vapor to escape during the drying process. A CentriVap connected to a cold trap is used to dry the pellet. Place the centrifuge tube inside the CentriVap, and evaporate the residual solvent for 30 minutes at 25°C. Once the drying cycle is complete, visually observe the pellet. It should be dry, white, and intact.

[0364] Reconstitution of protein pellet: Note that protein loading in the insert, along with the weight of the original dry insert sample, is used to determine how much buffer should be used to reconstitute the pellet. The pellet should be reconstituted to a concentration that is within the standard curve. Carefully remove the parafilm from the centrifuge tube, and reconstitute the pellet with calculated amount of 50 mM phosphate buffer, 1.2% Sodium chloride. Mix the solution by vortexing in 5 second intervals 2-3 times. Let the samples incubate at room temperature for 40 mins to ensure complete decomplexation of the protein from the leftover sheath polymer. Once the pellet is dissolved, there may be a minimal amount of white precipitate in the solution, e.g., if there is residual sheath polymer. Centrifuge the solution for 15 sec by selecting the short spin option to separate the buffer solution from the precipitate. Use a 200 pL pipet to take 100 pL of the clear phase, and place it inside a low-volume polypropylene HPLC vial. Assay the samples along with a fresh STD curve using the HPLC method for protein analysis. Examples 21 and 22 are examples of sandwich inserts, with protein of MW <20K and 60-70K.

[0365] When the sandwich insert is not used, the incorporated protein is almost immediately released. Fig. 39 demonstrates the necessity of the sandwich insert to achieve a sustained release profile of protein with % burst < 30%. Fig. 38 shows the appearance of core-shell co-axial insert, without a sandwich microstructure. The composition of the core-shell (middle) layer is identical to that of a sandwich insert. The insert in Fig. 38 shows electrospun fibers, not the unique microstructure shown in the sandwich insert (Fig. 41).

[0366] Fig. 39 shows the lack of sustained release from inserts that only contain the core-shell middle layer, without the encasing layers.

[0367] For core-sheath fibers with incorporated protein, the co-axial microstructure of the fibers was not adequate to provide controlled release of the bioactive compound. When encased in the form of the "sandwich," sustained release could be achieved.Example 24: Sandwich (Layered) Insert with Incorporated Protein (MW<25K) and Diffusion Aid

[0368] Table 61. AB-NANOM-X 07-51 Insert Composition

[0369] Table 62. Process Parameters, Insert Composition and AnalysisExample 25: Sandwich (Layered) Insert with Incorporated Protein (MW<70K) and Diffusion Aid

[0370] Table 63. AB-NANOM-03-56 Insert Composition

[0371] Table 64. Process Parameters, Insert Composition and AnalysisExample 26: Sandwich (Layered) Insert with Incorporated Protein (MW>100K) andDiffusion Aid PTMC(3k)-PEG(3.4k)-PTMC(3k)

[0372] Table 65. AB-NANOM-X-07-69 Composition

[0373] Table 66. Process Parameters, Insert Composition and AnalysisExample 27: Sandwich (Layered) Insert with Incorporated Protein (MW<70K) and Diffusion Aid ( containing PTMC homopolymer with an inherent viscosity of 0.65 dL / g)

[0374] Table 67. AB-NANOM-03-69 Insert Composition

[0375] Table 68. Process Parameters, Insert Composition and AnalysisExample 28: Sandwich (Layered) Insert with Incorporated Protein (MW<70K) and Diffusion Aid (containing PCL-PTMC (10:90) and PCL-PTMC (90:10) in outer layers)

[0376] Table 69. AB-NANOM-03-72 Composition

[0377] Table 70. Process Parameters, Insert Composition and Analysis

[0378] In-Vitro Release Profiles from "Sandwich" Inserts

[0379] Procedure: Punch out a 10 mm punch and record the weight. Take a polypropylene vial and record the empty tare weight of the vial. Place the insert inside the vial. Add 50 pL of phosphate buffer to the vial and record the weight. Place the vial in an incubator set to 37 °C. After 30 min, remove the vial from the incubator.Record the total weight of the vial. Remove 20-30 pL of the buffer and add it to the HPLC vial. Record the weight of the buffer removed. Replenish the 20-30 pL of buffer with fresh phosphate buffer with sorbitol and place the vials back in the incubator. Repeat the steps above for each time point (4 hours, 1 day, 2 days, etc.). Dilute the recovered buffer from every time point with 20 pL of phosphate buffer. Run the sample on the HPLC and analyze the amount of protein released using UPLC.

[0380] Fig. 40 demonstrates sustained release profiles from "sandwich" insert. Sustained release of the highly water-soluble proteins is not possible without the unique micro structure of the sandwich insert. The polymers that are able to flow are comprised of polymers that are poly (trimethylene carbonate), poly (trimethylene carbonate) copolymerized with other polymers such as poly(caprolactone), or blendedwith other polymers to impart strength. The molecular weight of the polytrimethylene carbonate would influence the flow properties of the encasing insert. For example, if the inherent viscosity of the poly(TMC)co-PCL is 1.2 dl / g, the encasing layer does partially form without annealing (30-35C) for a prolonged period, post-fabrication. Additionally, continuity is further established when the insert is subjected to sterilizing radiation at 6 kGy.

[0381] The "middle layer" remain as electrospun fibers (Fig. 42). As demonstrated in Fig. 41, the encasing layers also begin as electrospun fibers, but transform into continuous layers (top and bottom) as the encasing layer. The encasing layer polymer is necessary to slow down the diffusion of the incorporated bioactive agent. The role of the diffusion aid poly(trimethylene carbonate)-polyethylene oxide- poly(trimethylene carbonate) is to generate water soluble phases for the incorporated protein to diffuse through.

[0382] The polymers in the encasing layers need to be able to "flow" to transform the top layer into the continuous layer. These are polymers based on poly(trimethylene carbonate)-based polymers and blends and copolymers with pol(caprolactone) to add strength and durability.

[0383] In one example, when poly(trimethylene carbonate)-co-caprolactone (90:10) was used as the sole encasing polymer, the resultant product flowed too much, exposing the incorporated bioactive protein, resulting in -89% burst of protein. Thus, polycaprolactone is required to add strength and structure to the insert top and bottom layers. When fully encased, the protein burst is lowered to < 25%, but further release of protein is enabled by increasing the concentration of poly(TMC)-PEG-poly(TMC) in the encasing layers.

[0384] Effect of E- beam and temperature on flowability of the polymer.

[0385] This mesh has the film sheath composition as 18% w / v (90:10) PCL- PTMC: PCL with 0.116% PTMC-PEG-PTMC (diffusion aid). This composition had a good release profile (30 min burst 20.1%, post E-beam (6 kGy) 30 min burst 23.4%). Fig. 43 shows pre-annealing and post annealing SEM images. The image at T= 24 hr shows a drastic reduction in number of holes, which shows the effect of temperature (annealed at 33C) on the flowability of polymer. The third image shows the effect of E-beam on the structure of the mesh. The number of holes reduced slightly visually,but has a significant effect on the release profile. This is a good example showing the effect of temperature and E-beam on the release profile and the flowability of polymer.

[0386] The energy associated with ebeam radiation enables flow of the encasing polymer, having a similar effect as annealing. Fig. 44 demonstrates the rate of release of model proteins before and after annealing, with the % protein burst lowered from 60% (pre- annealing) to 35% (post- annealing). Scanning Electron Micrographs of the wafers post-annealing shows more a continuous outer layer. Fig. 45, on the other hand, shows release of a protein from inserts before and after e- beam sterilization. The release profile before sterilization releases rapidly with greater than 60% released in 1 day, whereas post-sterilization demonstrates a more sustained release of protein. As shown in Fig. 43, both annealing and energy from sterilizing radiation has the effect of causing flow of the outer layers to form diffusional barriers to protein release. Other sterilizing radiation such as X-Ray or gamma radiation would work similarly to e-beam (P-radiation). The sterilizing radiation would be in the range 6-25 kGy, preferably cold.

[0387] OTHER EMBODIMENTS

[0388] All of the features disclosed in this specification may be combined in any combination. Each feature disclosed in this specification may be replaced by an alternative feature serving the same, equivalent, or similar purpose. Thus, unless expressly stated otherwise, each feature disclosed is only an example of a generic series of equivalent or similar features.

[0389] From the above description, one skilled in the art can easily ascertain the essential characteristics of the present invention, and without departing from the spirit and scope thereof, can make various changes and modifications of the invention to adapt it to various usages and conditions. Thus, other embodiments are also within the scope of the following claims.

Claims

CLAIMSWhat is claimed is:

1. A dry drug delivery system comprising a middle layer encased in two outer layers, the middle layer and the two outer layers each comprising a mixture of electrospun fibers, the dry drug delivery system having a final combined density in the range 0.1-1 g / cm3, a thickness < 1 mm, an elastic modulus < 4 MPa, and a change in elastomeric strain >100%, wherein the two outer layers each include one or more polymers selected from ethylene vinyl acetate, poly trimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) having a copolymer ratio of 90:10, PCL, and P(TMC)-polyethylene oxide (PEG)-P(TMC) in which a molecular weight of the PEG segment is 1,000-20,000 Dal and a molecular weight of the P(TMC) segment is 1,000-10,000 Dal; and the middle layer being formed of core- shell fibers, each of which includes: a core that contains a bioactive substance and one or more excipients to protect the bioactive substance from degradation, and an outer shell formed of a hydrophobic elastomeric polymer comprised of P(TMC) and PCL in a weight ratio of 90: 10 to 30:70.

2. The dry drug delivery system of claim 1, wherein the ethylene vinyl acetate has a degree of vinyl acetate substitution of 42%, the P(TMC) has an inherent viscosity of 0.3- 1.2 dL / g, the P(TMC)-co-PCL has an inherent viscosity of 1.2- 1.6 dL / g) and a copolymer ratio of 90:10, and the PCL has an inherent viscosity of 1-1.5 dL / g).

3. The dry drug delivery system of claim 1 or claim 2, wherein the mixture of electrospun fibers in the two outer layers are annealed to form a continuous layer.

4. The dry drug delivery system of any one of claims 1 to 3, wherein a content of the one or more polymers in the two outer layers is 0-10% w / w P(TMC), 60-95% w / w P(TMC)-co-PCL, 5-20% w / w PCL, and 0-2.5% P(TMC)-PEG-P(TMC), based on the weight of the two outer layers.

5. The dry drug delivery system of any one of claims 1 to 4, wherein the system is biodegradable.

6. The dry drug delivery system of any one of claims 1 to 5, wherein the bioactive substance is a water-soluble drug selected from a small molecule of molecular weight < 1000 Dal, a large molecule of molecular weight > 1000 Dal, a protein of molecular weight between 10,000 Dal and 250,000 Dal, an amino acid, a peptide of 2-100 amino acid residues, or a nucleic acid.

7. The dry drug delivery system of claim 5, wherein the nucleic acid is a plasmid DNA, a linear DNA, an oligonucleotide, an mRNA, or an siRNA.

8. The dry drug delivery system of claim 1, wherein the bioactive substance is water-soluble and selected from the group consisting of an immunoglobulin, an antigen-binding domain, an anti-bacterial, an anti-viral, a steroid, an anti-glaucoma agent, an NSAID, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an anti-allergic, a lubricant, an antifungal, an anti-inflammatory, a protein, a peptide, a growth factor, an enzyme, a vitamin, a hormone, a polysaccharide, a peptide-nucleic acid, and a nucleic acid, or a mixture thereof.

9. The dry drug delivery system of claim 1, wherein the bioactive substance is hydrophobic and selected from an anti-microbial, a steroid, an NSAID, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an anti- allergic, a lubricant, an antifungal, an anti-inflammatory, and an anti-glaucoma agent.

10. The dry drug delivery system of any one of claims 1 to 9, further comprising water-soluble adhesive fibers deposited onto the system, the water-solubleadhesive fibers selected from polyacrylic acid, PEG-multi- armed amine, polyethyleneimine, polyamidoamine, polylysine, polyarginine, chitosan gluconate, and derivatives and blends thereof.

11. The dry drug delivery system of any one of claims 1 to 10, wherein the one or more excipients in the core of the core-shell fibers interact with the bioactive substance as a water substitute, complex with the bioactive substance, or form an interfacial layer between polymer-containing organic and water phases during fabrication of the system by electrospinning.

12. The dry drug delivery system of claim 11, wherein the one or more excipients are selected from the group consisting of hydroxypropyl P-cyclodextrin, gamma cyclodextrins, sulfobutyl P-cyclodextrins, PEG40- stearate, trehalose, sorbitol, sucrose, mannitol, poloxamer 407, poloxamer 188, polysorbate 80, polysorbate 20, dextran, polyvinyl alcohol, polyethylene oxides, polyvinyl pyrrolidone, hydroxypropyl methyl cellulose, human serum albumin, and soluble collagen, the one or more excipients being present in the dry drug delivery system in a concentration of 1-75% w / w.

13. The dry drug delivery system of claim 1, wherein the system has been sterilized by cold e-beam radiation at a dose range between 6-25 kGy.

14. The dry drug delivery system of claim 1, wherein the system is round, oval, crescent- shaped, elliptical, annular, square, rectangular, or tubular.

15. The dry drug delivery system of claim 1, wherein the hydrophobic elastomeric polymer forming the outer shell of the core-shell fibers is selected from the group consisting of ethylene vinyl acetate (EVA), polydioxanone, polysiloxane, P(TMC), and PCL, and the P(TMC)-co-PCL is present in the two outer layers at a concentration of 50-95% w / w based on the weight of the two outer layers.

16. A method for delivering a bioactive substance to a soft tissue of a subject, the method comprising:obtaining a dry drug delivery system of claim 1, and applying the system to a surface of a soft tissue of a subject before or after adding a hydrating fluid to the soft tissue, the hydrating fluid comprising an isotonic solution that contains one or more of sodium chloride, a phosphate, a citrate, albumin, a magnesium salt, a calcium salt, borate, boric acid, a balanced salt solution, multibranched PEG-amine, PVA, multi-branched polylysine, multi-branched polyarginine, guar gum, gellan, sodium alginate, xanthan gum, carboxymethyl cellulose, the hydrating fluid having a pH of 6-7.8 and an osmolality of 270-340 mOsm / kg.

17. A water-dissolvable dry device, the device comprising a physical mixture of a first fiber and a second fiber different from the first fiber, the first fiber containing sodium hyaluronate, polyethylene glycol (PEG), and polysorbate 80 or polysorbate 20, and, optionally, castor oil, a buffer salt, an amino acid, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinyl pyrrolidone, dextran 70, albumin, PEG-multi-arm amine, collagen type 1, glycerol, or PEG40-stearate, and the second fiber containing polyvinyl alcohol (PVA) and sodium alginate and, optionally, sodium borate, dextran 70, PEG, PEG-multi-arm- succinimidyl ester, polypropylene glycol, propylene glycol, collagen type I, a bioactive substance, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or beta 1>6 glucan, wherein the first fiber and the second fiber are configured to react with each other and with a biological tissue upon hydration, and the device transform into a hydrogel upon contact with the biological tissue and subsequent hydration and completely dissolves over a period of 0 h to 14 days after contacting the biological tissue.

18. The water-dissolvable dry device of claim 17, wherein the PVA has a molecular weight of 88,000-130,000 Dal with degrees of hydrolysis between 88- 100%, the sodium hyaluronate has a molecular weight between 50,000-5,000,000 Dal and the sodium alginate has a molecular weight between 32,000-400,000 Dal, the first fiber including 80-90% by weight PEG, 0.5-5% by weight sodium hyaluronate, 0.1- 0.3% by weight sodium chloride, 0.001-2% by weight polysorbate 20 or polysorbate 80, and the second fiber including 80-95% by weight PVA.

19. The water-dissolvable dry device of claim 17 or 18, further comprising a bioactive substance selected from a small molecule of molecular weight < 1000 Dal, a large molecule of molecular weight > 1000 Dal, a protein of molecular weight between 10,000 Dal and 250,000 Dal, an amino acid, a peptide of 2-100 amino acid residues, or a nucleic acid.

20. The water-dissolvable dry device of claim 19, wherein the bioactive substance is water-soluble, minimally water-soluble, or water-insoluble.

21. The water-dissolvable dry device of claim 17, wherein, upon placement on a biological tissue, the device dissolves by simple solubilization in aqueous fluids or by biodegradation.

22. The water-dissolvable dry device of any one of claims 17 to 21, wherein the device is round, oval, crescent-shaped, elliptical, annular, square, rectangular, or tubular.

23. The water-dissolvable dry device of claim 17, wherein the device has been sterilized by cold e-beam radiation at a dose range between 6-25 kGy.

24. The water-dissolvable dry device of claim 19, wherein the first fiber or the second fiber includes the bioactive substance, the bioactive substance being selected from an immunoglobulin, an antigen-binding domain, an anti-bacterial, an anti-viral, a steroid, an anti-glaucoma agent, an NSAID, an UV blocking agent, a healing agent, a decongestant, an antihistamine, an antioxidant, a vasoconstrictor, an anti-allergic, a lubricant, an analgesic, an antifungal, an anti-inflammatory, a protein, a peptide, a growth factor, an enzyme, a vitamin, a hormone, a polysaccharide, a peptide-nucleic acid, a nucleic acid, or a mixture thereof.

25. The water-dissolvable dry device of claim 21, wherein the biological tissue is the ocular mucosa, vaginal mucosa, rectal mucosa, nasal mucosa, or oral mucosa.

26. A method for delivering a bioactive substance to a soft tissue of a subject, the method comprising: obtaining a water-dissolvable dry device of claim 17, and applying the system to a surface of a soft tissue of a subject, wherein the device transform into a hydrogel upon contact with the surface of the soft tissue upon hydration.