Methods and compositions for applying a therapeutic agent to a soft tissue surface

Abstract

A dry drug delivery system is disclosed, the system comprising an intermediate layer encased in two outer layers, the intermediate 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 first fibers and second fibers, wherein the first fibers and the second fibers are configured to react with each other and with biological tissue upon hydration, and the device transitions to a hydrogel upon contact with the biological tissue. Also provided is a method for delivering a bioactive substance to soft tissue using the system or device.

Description

[0001] Related applications

[0002] This application claims the benefit of U.S. Provisional Application Serial No. 63 / 551,400, filed February 8, 2024. The full teachings of the foregoing application are incorporated herein by reference. Background Technology

[0003] Limited Challenge

[0004] Problems with unstable molecules in compounds include low encapsulation efficiency and degradation, irritation, low tissue retention, and rapid release kinetics.

[0005] Low encapsulation efficiency due to degradation of unstable drugs: Highly water-soluble small molecules (synthetic and biological) and larger molecules (such as peptides, proteins and nucleic acids) are difficult to formulate in matrices that can encapsulate high concentrations of unstable drugs (such as proteins or DNA) and achieve a sustained release pattern of the intact drug.

[0006] These molecules (both small and large) can be unstable and readily degraded under specific microenvironments (pH, salt, enzymes, organic / aqueous interfaces, solvents), environmental conditions (high temperatures, etc.), physical shear forces, ionic interactions, and exposure to extreme conditions. Strategies for constructing these mechanisms involve selecting optimal pH and buffer counterions, and screening excipients and excipient blends that help stabilize these unstable molecules in both solution and solid forms. While formulations using specific excipients can protect unstable molecules from proteases and microenvironmental stresses, achieving high concentrations is extremely difficult.

[0007] Furthermore, for localized tissue surfaces, sustained-release formulations present the challenge of retention at the target site. For localized tissue surfaces (e.g., skin, sublingual, intranasal, vaginal, and ocular), sustained-release “film” formulations are a preferred strategy. For tissue spaces such as the ocular mucosa, film formulations are inserts. In cavitary tissues (e.g., intranasal, sublingual, rectal, and buccal tissues), sustained drug delivery is challenging due to rapid fluid flow. Another tissue type is the esophageal mucosa, which requires in-situ retention for site-specific sustained drug administration. Vaginal, rectal, intranasal, and urinary tract mucosa can benefit from sustained release from inserts, especially where gravity may displace any viscous solutions, gels, and ointments. Inserts that adhere well to tissues (mucosa and non-mucosa) and resist displacement due to gravity enable site-specific targeted drug delivery without wasting drug.

[0008] Incorporating highly water-soluble compounds into polymeric membrane implants in sustained-release form (with the compound dispersed within a polymer matrix) is challenging. For implants less than 1 mm thick, the encapsulated compound exhibits a short diffusion path length, leading to rapid drug release. Due to the small diffusion path length, rapid release kinetics were observed for both hydrophobic and hydrophilic molecules.

[0009] Another challenge arises when unstable active pharmaceutical ingredients (APIs) or mixtures of pharmaceutical ingredients need to be encapsulated in their intact form within a polymer delivery system and must remain intact, or maintain a high degree of integrity during storage and release from the drug delivery system matrix. The components of the delivery system cannot react with the active ingredient, which could impair its bioavailability. These unstable molecules typically fall into the category of biomolecules such as antibodies, proteins, peptides, aptamers, and polynucleotides such as DNA. Additionally, compounds that are readily degraded through hydrolysis, oxidation, or other mechanisms can benefit from encapsulation in a matrix composition that not only sustains their release but also remains stable against degradation factors such as proteases, nucleases, oxidants, hydrolysants, and other degradative substances. Therefore, all components of the delivery system must be chemically compatible with the encapsulated compound and act as a protector to maintain its integrity.

[0010] Finally, the membrane implant placed on the target tissue should be removable after the drug is depleted from the matrix. The implant placed on the tissue should also be soluble, i.e., dissolve after drug release or, in the case of a "drug-free" implant, after placement. The implant should be tissue-biocompatible, and the dissolution of its components should be modulated.

[0011] Irritating issues

[0012] A fundamental challenge inherent in implantable and localized drug delivery systems is irritation caused by undesirable molecular interactions between the implant and underlying tissue, particularly in tissues containing moisture, mucous membranes, and nerve endings. Irritation caused by implants at the local tissue site is a common foreign body reaction. One of the most common problems is that implants placed on mucous membranes cause unpleasant irritation, as well as congestion, local discomfort, pain, and erythema. However, some of the biomaterials that make up the implant (PLGA, PLA, PCL, and copolymers and blends) harden at physiological temperatures and pH, shrinking into devices with sharp edges and losing their elasticity and flexibility. Hard, non-flexible implants become highly irritating to underlying tissues, especially in the eye, where previously flexible implants can feel like shards of glass within hours. PLGA degrades into acidic fragments, resulting in a dense surface accumulation of acidic impurities in the local microenvironment of the implanted implant. However, even transient acidic microenvironments can lead to biocompatibility issues, including immune responses manifesting as mild irritation and congestion, or more severe, such as burning sensations. Due to their enhanced cellular uptake and biodistribution, PLGA particles are known to affect cell growth, activity, and tissue response, and clinically manifest as inflammation, irritation, and congestion. These issues necessitate the continued development of innovative approaches to create biocompatible tissue-implant interfaces that minimize irritation.

[0013] The issue of implantation stationing

[0014] Mucous membranes present challenges for implant retention in both accessible and inaccessible spaces. These include the mucous membranes of the eyes, nose, rectum, vagina and cervix, uterus, urinary tract, esophagus, cheeks, sublingual region, lungs, and bronchi.

[0015] One of the key performance criteria for drug delivery systems applied to tissue surfaces is the long-term residence of the drug-containing device. "Long-term" can be relative: any duration between a few hours and many days (e.g., 5–180 days). Electrospun implants placed in the conjunctival fornix (lower and upper eyelids) are characterized by their high interstitial porosity, allowing the biological environment to circulate and thus integrating the implant into the underlying tissue. Conversely, PLGA-based electrospun or solvent-cast implants lose shape and undergo dimensional shrinkage. The combination of dimensional instability and increased stiffness leads to device delamination from the underlying tissue, causing localized irritation due to friction of sharp edges against soft tissue, and inevitably resulting in detachment from the application site. Summary of the Invention

[0016] To overcome the above challenges, a dry drug delivery system is provided, comprising an intermediate layer enclosed within two outer layers, the intermediate layer and the two outer layers each comprising a mixture of electrospun fibers, the two outer layers each comprising one or more polymers selected from the group consisting of ethyl vinyl acetate, polytrimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) in a copolymer ratio of 90:10, PCL, and P(TMC)-polyethylene oxide (PEG)-P(TMC), wherein the intermediate layer is formed of core-sheath type fibers, each core-sheath type fiber comprising: (i) a core containing a bioactive substance and one or more excipients protecting the bioactive substance from degradation, and (ii) a sheath formed of a hydrophobic elastic polymer.

[0017] A method is also provided for delivering a bioactive substance to a soft tissue by applying the dry drug delivery system to the surface of the soft tissue before or after adding a hydration fluid to the soft tissue of a subject.

[0018] Additionally, a water-soluble drying device is disclosed, which is a physical mixture of a first fiber and a second fiber, wherein the first fiber contains one or more of sodium hyaluronate, castor oil, polyvinyl alcohol (PVA), and polysorbate 80 or polysorbate 20, and optionally contains buffer salts, amino acids, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinylpyrrolidone, dextran 70, albumin, PEG-polyarm amine, polyethylene glycol, type I collagen, glycerol, and PEG40-stearate, and the second fiber contains PVA and sodium alginate, and optionally contains sodium borate, dextran 70, PEG-polyarm succinimide, polypropylene glycol, propylene glycol, type I collagen, bioactive substances, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or β1>6 glucan. The first fiber and the second fiber react with each other and with biological tissue when hydrated, and the device transforms into a hydrogel when in contact with the biological tissue.

[0019] Details of one or more embodiments of the invention are illustrated in the following description and accompanying drawings. Other features, objects, and advantages of the invention will become apparent from the description, drawings, and claims. Attached Figure Description

[0020] The foregoing and other objects, features and advantages will become apparent from the following description of specific embodiments of the present disclosure as shown in the accompanying drawings.

[0021] Figure 1Exemplary size and shape of an ophthalmic implant

[0022] Figure 2 Exemplary application tool design

[0023] Figure 3 (a) Top left: Scanning electron microscope image of the soluble implant; (b) Bottom left: Scanning electron microscope image of the ocular mucosa; (c) Hydrated implant placed on the ocular surface.

[0024] Figure 4 Young's modulus of dry insert

[0025] Figure 5 Young's modulus of the hydrated insert

[0026] Figure 6 Scanning electron microscope images of soluble in-situ crosslinkable implants (PVA-alginate).

[0027] Figure 7 Crosslinking between borate ions and hydroxyl groups in the polymer chain of polyvinyl alcohol.

[0028] Figure 8 Crosslinking kinetics and solubility of collagen-containing hydrogel formulations

[0029] Figure 9 Hydrogel implantation

[0030] Figure 10 Hydrogel formation

[0031] Figure 11 The implant is applied to the conjunctival fornix.

[0032] Figure 12 Application of wetting solution

[0033] Figure 13 : The hydrogel placement component after wetting

[0034] Figure 14 The implant was inserted into the tissue in both eyes one hour later.

[0035] Figure 15 Hydrogel implantation

[0036] Figure 16 Hydrogel was not visible 24 hours after implantation.

[0037] Figure 17 After the implant is applied

[0038] Figure 18 Hydrogel formation after rapid water absorption

[0039] Figure 19 : On the surface of the eye before application

[0040] Figure 20 SB-NANOM-X-15-14 Insertion Application

[0041] Figure 21 Postmortem ocular surface, T=6 hours

[0042] Figure 22 Bioadhesion occurring ~30 minutes after application of the implant and hydration buffer (Left = implant peel strength from tissue at T=0; Right = implant peel strength from tissue at T=30 minutes).

[0043] Figure 23 Peptides before and after incorporation into the implant.

[0044] Figure 24 : Annular hydrogel sheath with hydrophobic core fiber for mucosal adhesion

[0045] Figure 25 Release of the small peptide exenatide from the cross-linked, soluble implant

[0046] Figure 26 Small peptide release from electrospinning insert

[0047] Figure 27 The implant exhibits high elastomer properties.

[0048] Figure 28 Drug products containing IgG

[0049] Figure 29 Release of intact proteins from implanted drug products

[0050] Figure 30 Sustained release of IgG from implanted drug products containing P(TMC)-lactide

[0051] Figure 31 Batch reproducibility of products containing IgG implants

[0052] Figure 32 Size exclusion chromatograms of protein-containing implants sterilized by electron beam exposure at 25 kGy (left) and 15 kGy (right).

[0053] Figure 33 Comparison of dimensional stability: PCL12 (blend), PLGA, PDO and PCL80K

[0054] Figure 34The PCL12 blend composition at t0 (left) and t14 days (right) shows dimensional stability over time.

[0055] Figure 35 t0 (left) and t9 (right) show the dimensional stability over time.

[0056] Figure 36: (A) A 6 mm insert to be placed on the tissue; (B) A square insert to be placed on the tissue; (C, D) An 8 mm insert.

[0057] Figure 37 Image A: Score = 0 (top left), no adhesion shown and scored 0. Image B: Score = 1 (top right), shows some adhesion to the tissue, peeling, and stretching of the conjunctival tissue. Image C: Score = 2 (bottom left), greater resistance to detachment, the mesh begins to stretch and tear. Image D: Score = 3 (bottom right), greater adhesion, the sample can be lifted through the mesh.

[0058] Figure 38 Scanning electron microscope image of a core-skin type implant.

[0059] Figure 39 In vitro release of proteins (core-sheath type vs. sandwich type)

[0060] Figure 40 : Sustained release mode of proteins from the "sandwich" insert

[0061] Figure 41 Scanning electron microscope images of sandwich inserts before annealing, after annealing, and in optimized annealing states.

[0062] Figure 42 Scanning electron microscope image of a partially annealed sandwich insert (expansion, revealing core-shell type fibers).

[0063] Figure 43 The effect of annealing and electron beam sterilization on the microstructure of the implant (AB-NANOM-X-07-65).

[0064] Figure 44 Effect of annealing on the release rate of model proteins (MW 85-140K)

[0065] Figure 45 Effects of electron beam sterilization on the in vitro release of bound proteins (MW <25K). Detailed Implementation

[0066] This invention describes a variety of different strategies for delivering therapeutic molecules (proteins, peptides, nucleic acids, or small molecules) via electrospinning (a) "layered" (sandwich-type) implants, while also providing means for regulating the release of bioactive substances, protecting the integrity of the encapsulating material, improving the biocompatibility of the device with the underlying tissue, and the persistence of the device at the application site; and (b) soluble implants, wherein the implant dissolves into a thin, transparent film on the tissue surface after placement, utilizing incorporated components (whether drug-containing or drug-free) to form an air-oil-water interface on the tissue surface, aiming to alleviate symptoms on the tissue surface.

[0067] Sandwich-type (layered) insert:

[0068] The compositions of fiber implants prepared by the electrospinning and electrospraying processes described herein focus on: (a) encapsulating concentrated molecules with different levels of water solubility and instability (stability); (b) modifying the composition and method to release the molecules in a sustained manner; (c) minimizing irritation at the local tissue site; and (d) maximizing retention at the local delivery site.

[0069] The described implant has a hydrophobic drug release barrier in the form of a hydrophobic outer "skin" layer covering a hydrophilic "core" layer containing the drug. To achieve a linear release rate, "sandwich" implants have been described in which drug-containing fibers (core-skin type or integral type) are sandwiched between two polymer layers ("coating layers").

[0070] The sandwich device has three layers. The two outer layers (coating layers) have a hydrophobic polymer composition that modulates the release of the encapsulating material, provides flexibility, exhibits tissue compliance (wraps around the tissue), and possesses dimensional strength and stability upon hydration by tissue fluid. The two outer layers are formed by electrospinning, and they form fibers upon deposition but "flow" upon drying, thus forming a continuous or nearly continuous membrane covering / coating the inner layer.

[0071] The inner layer of the sandwich device is composed of fibers (which remain in fibrous form, unlike the coating layer that transforms into a continuous coating layer during drying). Depending on the desired release mode, the inner layer may contain single-component fibers (single-channel) or core-shell type fibers (dual-channel).

[0072] The coating polymers possess properties that allow them to "flow" and form coating layers. This "flow" can be achieved if the glass transition temperature of the coating polymer is within acceptable limits.

[0073] The coating layer can be selected from the group derived from poly(trimethylene carbonate) (PTMC). The derivative can be PTMC derivatized using polycaprolactone, poly(lactide), or poly(siloxane), thereby imparting various properties to the final polymer. PTMC of one molecular weight can be blended with PTMC of another molecular weight to provide flow properties for forming continuous or nearly continuous sandwich layers.

[0074] The purpose of developing layered drug delivery systems is to use elastomeric polymers to create flexible tissue-device interfaces to eliminate or minimize irritation. If the modulus of the tissue matches the modulus of the implant, or if the modulus of the implant is less than that of the tissue, the implant will not delaminate or slide across the tissue, causing irritation. The properties of the polymer coating layers enable tissue conformation, which improves retention on mucosal surfaces.

[0075] The inner coating contains a drug that can be released in a sustained manner. To achieve this, the drug may be contained within a core-shell type fibrous matrix. Optionally, the drug may also be contained within the coating layer.

[0076] The implant can be a ring-shaped (“ring”), elliptical, curved, or tubular implant (film). The tube can be 1-10 mm in length and 0.2-10 mm in diameter.

[0077] The drugs discussed can be large molecules or small molecules, or combinations of multiple molecules. Molecules can be of biological origin or synthetic. Molecules can be proteins, portions of proteins, nucleic acids, or peptides, or peptide-nucleic acid combinations, or small molecules with varying degrees of water solubility and stability.

[0078] Due to the "thin" nature of the implant, the delivery device is biodegradable and degrades through surface erosion.

[0079] All layers of the implant are prepared using a fiber electrospinning process, with the fibers deposited onto a collector. However, most of the polymer in the electrospinning remains in fibrous form. To form the coating layer, PTMC polymers, PTMC derivatives, and blends are required. These types of layered drug delivery systems can be used to deliver water-soluble and water-insoluble compounds. These compounds can be small molecules or large molecular biological agents, such as antibodies, nucleic acids, and inactivated viruses.

[0080] Characteristics of the inner layer of a sandwich device

[0081] A drug delivery system (implant) features a drug encapsulated within an internal "core" (composed of a water-soluble protective component) and coated with a hydrophobic polymer used to create a "skin," thereby enabling slow release from the matrix. For unstable drug molecules within the "core," the polymer "skin" matrix provides protection against degradation factors and acts as a barrier to drug release. The core components protect the drug from degradation (e.g., oxidation, proteolysis, deamidation).

[0082] Additional outer coatings can be deposited onto the implant, "sandwiching" core-shell fibers by coating the fibers with adjacent layers. This "sandwiching" or coating is used to modulate drug release and provide a flexible, biocompatible surface for contact with tissues.

[0083] Encapsulation of unstable water-soluble molecules in the inner "core" of a sandwich device.

[0084] The drug delivery system (implant) contains a drug encapsulated in a core within a hydrophobic "skin" polymer to achieve slow release from the matrix while being protected from endogenous proteases, nucleases, or oxidizing agents known to cause degradation. For drug protection, excipients may be included to protect the drug from oxidation, deamidation, and aggregation.

[0085] To achieve sustained release of highly water-soluble drugs

[0086] Water-soluble bioactive molecules are very difficult to encapsulate at concentrations that allow for sustained drug release. One way to effectively encapsulate water-soluble proteins in a hydrophobic polymer matrix is ​​to create a core-skin type placement having a protein as a hydrophilic core and a ring-shaped polymer coating as a hydrophobic barrier for release. To achieve this, the water-soluble drug, such as the protein, needs to be dissolved in an aqueous medium, and the hydrophobic polymer needs to be dissolved in an organic solvent. Methods for minimizing protein precipitation at the water-organic interface are disclosed below. Methods for creating interfaces that are biocompatible with the components present in both phases are also disclosed.

[0087] Another effective way to encapsulate drugs is by sandwiching drug-containing electrospun fibers between two consecutive membrane layers, thus effectively encapsulating the drug in a fully encapsulated drug delivery system. The highly elastic outer layer is used to "shape" itself onto the tissue surface, thereby minimizing or eliminating any biomaterial-tissue interaction. Sandwiching drug-containing electrospun fibers between two elastic hydrophobic membrane layers allows for sustained release over the desired time. The membrane sandwich layer is a polymer with flow properties after drying; polytrimethylene carbonate and its derivatives possess properties that enable membrane formation. The sandwich-type placement structure allows drug molecules to be encapsulated in the "middle" layer, while the outer layer acts as a release barrier.

[0088] Use of polymers with dimensional and modulus stability

[0089] The selection of materials that do not harden (i.e., modulus stability) or shrink (dimensional stability) when exposed to a biological environment is a key criterion of this invention. Certain biodegradable polymers (PLGA and PLGA copolymers) used in implants become hard and inflexible (modulus change) and shrink (dimensional change) after hydration in buffer solutions at physiological temperatures (37ºC) and pH (7.2-7.4). This causes the implant to become a dimensionally reduced fragment with sharp edges, leading to delamination from the tissue surface and irritation due to loss of flexibility. In contrast, implants constructed as blends of amorphous polymers (such as polycaprolactone, polytrimethylene carbonate, and poly(trimethylene carbonate: polycaprolactone), ethylene vinyl acetate polymers, polysiloxanes, and polydioxanone) exhibit dimensional stability and also possess sufficient elasticity to flexibly conform to tissue surfaces. By employing different ratios of random copolymers of polytrimethylene carbonate and polycaprolactone, implants with varying strengths and elastic properties can be achieved. Blends of ethylene vinyl acetate (EVA) and polycaprolactone exhibit dimensional stability and elastic properties. Copolymers or blends of various combinations of polytrimethylene carbonate, polycaprolactone, polydioxanone, ethylene-vinyl acetate, and polysiloxanes are suitable for this purpose.

[0090] This invention describes compositions and methods for formulating drug-containing implants, which are prepared by electrospinning, electrospraying, or a combination of electrospinning and electrospraying, with the aim of applying them to tissue surfaces or implanting them into tissues. The objectives of this invention are to improve the biocompatibility of the biomaterial-tissue interface by minimizing irritation, to enhance implant retention within tissues by matching the flexural and compressive moduli of the tissue to the polymer implant material, and to regulate the release of encapsulated drugs by modulating the matrix polymer composition and microstructure of the encapsulating polymer layer.

[0091] The implant can be placed within a tissue space as an implant, thus surgically placing the implant into the tissue space. For subcutaneous spaces, a drug-containing implant can be surgically placed into the tissue space. The implant can also be implanted into tissues that may contain a large amount of fluid (such as urinary tract spaces, such as in the bladder). In another embodiment, a local implant can be placed in a sinusoidal cavity. In another embodiment, a continuously drug-eluting implant can be placed in other tissue surfaces (such as rectal tissue, vaginal tissue, and esophageal tissue). The same or similar considerations apply to other tissues that can benefit from a drug-delivery implant (such as intranasal, intrabladder, sublingual, intragingival, buccal, intravaginal, intraesophageal, intracerebral, and intrarectal tissues).

[0092] After adhesion to tissue, the implant, which is bound to the tissue (via ionic or hydrogen bonds), can be removed by applying an isotonic salt / excipient composition. A composition of a hydrating solution that will cause the implant to detach from the mucosal tissue surface is disclosed.

[0093] In one embodiment, the implant design incorporates a hydrogel coating onto each implant fiber that adheres to the tissue surface, while also enhancing biocompatibility due to its high water content. The implant fiber layer contains a drug within its "core." The drug can be hydrophobic, hydrophilic, or amphiphilic. The hydrogel coating acts as a "sheath" encapsulating each fiber.

[0094] The adhesion of the implant is enhanced by applying an hydrating fluid (wetting solution), resulting in adhesion within 1-30 minutes.

[0095] In another embodiment, the implant comprises an inner hydrophilic core (which also contains the drug) and an outer hydrophobic flexible elastic sheath, and is co-spun with a biocompatible adhesive. The implant can be removed by an application fluid and administered using an application device specifically designed for this purpose. The implant is sterilized under specific sterilization conditions. In another embodiment, the drug delivery system does not use a bioadhesive, but is held in place by tissue geometry and / or tissue folds. Additionally, methods and compositions for encapsulating and sustainably releasing intact, highly water-soluble bioactive compounds (such as proteins, peptides, nucleic acids, and small molecules) in electrospun nanofiber implants are disclosed.

[0096] Improve biocompatibility and retention by coating each fiber of the implant.

[0097] One strategy to improve biocompatibility and retention is through the creation of nanoporous mesh implants containing hydrophobic biodegradable fibers encapsulated in hydrophilic polymers that react with each other and with tissue proteins, forming amide covalent bonds. This anchors the drug delivery system to the tissue site and encapsulates the hydrophobic polymer fibers. In the case of PLGA, which degrades through bulk erosion, encapsulation with a hydrophilic sheath prevents the formation of particulate debris in local tissue spaces, thus ensuring long-term biocompatibility of the drug delivery system. The composition of the encapsulation focuses on branched polyethylene oxide polymers with reactive functional groups designed to react not only with each other but also with tissue proteins. The presence of polyethylene oxide promotes additional biocompatibility due to its high water content. Therefore, the body's immune regulatory cells "see" the highly hydrophilic surface and neutral pH. The encapsulation also contains polyvinyl alcohol, a filler molecule capable of modulating the crosslinking density of the hydrophilic ring sheath.

[0098] The second type of electrospun fiber implant composition is a soluble implant that transforms into a hydrogel upon placement on tissue and optionally cross-links with itself and the tissue. It slowly dissolves over time in a hydrated tissue environment, serving as a means of lubricating the eye or releasing compounds (which may be drug molecules). The soluble implant may initially be opaque but gradually becomes transparent or translucent as it absorbs water. The dissolution time will range from immediately after placement to 24 hours after placement.

[0099] The term "soluble" implies slow or rapid dissolution of the implant within the tissue space in which it is placed. All components of a soluble implant dissolve through dissolution or emulsification. The implant is soluble, tissue-conforming, highly hydrated, and adheres to the tissue. Soluble implants can be used in all mucosal tissues.

[0100] The focus of drug delivery systems is: (a) to produce biocompatible, highly flexible electrospun surfaces that cause minimal or no irritation to the biomaterial-tissue interface; (b) to produce implants that enable sustained release of proteins, nucleic acids, and small molecules; (c) to produce implants that encapsulate high concentrations of proteins without causing protein aggregation; and (d) to produce implant compositions that can be sterilized by various sterilization methods without losing the integrity of the encapsulated proteins.

[0101] The "sandwich" implant has three layers: (a) a lower hydrophobic elastic layer, (b) an intermediate layer containing the therapeutic agent to be delivered, and (c) an upper hydrophobic elastic layer. Other variations of the "sandwich" implant would involve a simpler intermediate layer containing a polymer of any drug, which is either hydrophilic (MW>100) or hydrophobic (MW>100), synthetic or biologically derived. Therefore, the intermediate layer can contain any type of drug. The polymer constituting the implant can be non-ionic or non-charged, but rather elastic, to generate van der Waals forces with the tissue surface, resulting in long-term adhesion.

[0102] The tissue type can be skin, eye, vagina, intranasal cavity, rectum, esophagus, urinary tract, sublingual cavity, or buccal cavity. The resulting device can be a circular disc, an oval disc, a shape specifically matched to the conjunctival fornix, or a ring-shaped device matching the ocular surface. For ocular use, the circular disc can have a diameter of 2 mm to 10 mm, preferably 3 mm to 6 mm, and most preferably 4 mm to 6 mm. For skin use, the device diameter can be 10-50 mm. For intranasal use, the device diameter can be less than 10 mm. For urinary tract use, the device can be inserted into the bladder with the assistance of a catheter. Devices for urinary tract drug delivery can be tubular in structure.

[0103] Characteristics of the coating layer

[0104] The top and bottom layers will contain elastic polymers such as poly(TMC), poly(TMC-polycaprolactone), poly(caprolactone), polydioxanone, ethylene-vinyl acetate, and blends thereof to achieve the property of producing a continuous film layer by electrospinning. In a variation, the top and bottom layers will contain PEGylated compounds (diffusion aids) to regulate the release of the drug from the intermediate layer.

[0105] The implant has an elastic elongation of >100%.

[0106] The implant has an elastic elongation of >100% to conform to the tissue. To test the mechanical adhesion strength of the implant to the tissue, stress-strain curves are generated to determine the fracture strain. Figure 27 The insertion component exhibits an elongation of >100%. This test was conducted using a Shimadzu mechanical testing machine.

[0107] Experiments were conducted using fresh bovine conjunctiva to compare implants with different percentages of elastic elongation. Implants with elongation (% strain) <100% were not flexible and did not conform to the underlying tissue. In contrast, implants with elongation >100% conformed to the underlying tissue.

[0108] Characteristics of the (intermediate) layer

[0109] The intermediate layer has a core-sheath structure and optionally has an integral layer (not a core-sheath type).

[0110] The purpose of the intermediate layer is to encapsulate hydrophilic compounds, thereby enabling: (a) high concentrations of compounds (proteins, etc.) to be encapsulated stably, and (b) excipients that stabilize proteins to remain encapsulated along with the proteins. Water-soluble drugs are contained within the core of the annular fiber structure of the electrospun insert, while the sheath contains an encapsulating hydrophobic polymer. The sheath's ability to act as an adjacent layer limits its ability to slow the release of the encapsulated compound, although the top / bottom layers of the insert primarily serve this function.

[0111] The microstructure of the implant allows unstable, water-soluble molecules to be encapsulated within the core-skin type fibers of the implant, thereby achieving protection against proteases, nucleases, other enzymes, and oxidative degradation. The drug delivery system (implant) contains a drug encapsulated within a core by a hydrophobic "skin" polymer to slow release from the matrix while being protected against endogenous proteases, nucleases, or oxidizing agents known to cause degradation.

[0112] For drug protection, excipients can be included to protect the drug from oxidation, deamidation, etc. The outer layer is hydrophobic and is an encapsulating polymer, ideally creating a continuous layer. The drug molecule encapsulated in the "core" can be hydrophilic, hydrophobic, or amphiphilic. Drug molecules can also be hydrophobic, in the form of free bases or free acids, but become water-soluble by converting to salt form.

[0113] The salt form may be selected from those available in the list of drug salt forms, including but not limited to acetate, hydrochloride, malate, iodide, bromide, chloride, malonate, tartrate, gluconate, glycolate, or any other salt form that can dissolve a water-insoluble drug in water. The active molecule may be amphiphilic, i.e., soluble in both aqueous and non-aqueous media. The active molecule may be selectively soluble with varying pH, ranging from 4.2 to 7.4, with a preferred pH range of 5.5-6.5 and another preferred pH range of 4.2-5.5.

[0114] Active drug molecules can be biological agents, such as antioxidants, peptides, peptide-nucleic acids, plasmid DNA or linear DNA, and proteins. Drugs can be of biological origin or synthetic, or synthetic derivatives of biomolecules. Preferred molecules are proteins, with a preferred molecular weight of 10-30 kDa, a preferred molecular weight of 30 kDa to 100 kDa, a preferred molecular weight of 100 kDa to 150 kDa, and a preferred molecular weight of 150-250 kDa. Molecules can be nucleic acids, such as antisense oligonucleotides, messenger RNA, RNAi, aptamers, or peptide-nucleic acid (PNA). Molecules can be small or large molecules, water-soluble molecules, or molecules that have become less water-soluble through salt formation or derivatization with other molecules. Molecules can be antiglaucoma agents, anti-inflammatory drugs, or antimicrobial agents. Molecules can be muscarinic agents or agents with therapeutic effects. Active agents can be combined with other molecules to achieve better encapsulation. The active drugs contained in the "core" can be formulated into nano-forms, such as calcium complexes of DNA, or complexes of positively charged peptides and negatively charged polymers (such as hyaluronic acid and xanthan gum).

[0115] The drug-containing solution may contain a dissolving excipient that reduces surface tension, thereby enabling the electrospinning process. The surface tension of the aqueous liquid (containing a water-soluble active ingredient) is in the range of 25-55 mN / m, measured by an Attension™ θ-contact angle optical tensiometer. Preferably, the surface tension of the aqueous liquid is in the range of 20-45 mN / m. In a more preferred range, the surface tension of the aqueous drug-containing core is in the range of 30-40 mN / m, and most preferably in the range of 30-36 mN / m. The conductivity of the aqueous core is in the range of 2000-6000 µS / cm, and preferably in the range of 2000-4000 µS / cm. The viscosity of the aqueous core solution containing the drug (peptide, protein, small molecule, nucleic acid) is in the range of 1-900 cP (25ºC, shear rate 40 s⁻¹), and the viscosity of the encapsulation layer is in the range of 1500-10,000 cP, preferably in the range of 1500-6000 cP. The viscosity was measured using an Anton Paar MCR92 cone-plate rheometer with a rotor CP50-0.5 / T. The interfacial tension between the aqueous and organic phases is less than 30 mN / m, preferably <20 mN / m. Electrospinning of the core solution and the encapsulating polymer solution is performed using a triaxial apparatus, wherein the organic vapor flows 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 of 15-21 cm, preferably in the range of 16-19 cm. The temperature of the processing chamber is in the range of 15-25ºC, and the relative humidity (RH) in the processing chamber is in the range of 25-40% RH, preferably less than 25-35% RH, and most preferably 25-30% RH. The encapsulating polymer solution can be at room temperature or slightly heated, within an acceptable range of 25-50ºC.

[0116] Aqueous core solutions containing water-soluble drugs may contain excipients that protect the active ingredient. These excipients include, but are not limited to, histidine, glycine, methionine, sorbitol, sucrose, trehalose, sodium chloride, glycerol, hydroxypropyl β-cyclodextrin, cyclodextrin, polyethylene glycol 400, polyethylene glycol 8000-200,000, propylene glycol, polypropylene glycol 400-5,000, polyvinyl alcohol, polyvinylpyrrolidone, PEG2K-DSPE, PEG5K-DSPE, PEG2K-DPPC, PEGylated phosphatidylcholine, phosphatidic acid, and lecithin. Aqueous solutions may contain proteins as excipients, such as collagen or albumin.

[0117] An interface is created between the aqueous and organic phases using excipients.

[0118] Proteins are typically unstable at the water-organic interface, where denaturation and loss of tertiary structure are common. The working hypothesis is that an amphiphilic surfactant monolayer at the interface would protect proteins from immediate denaturation events leading to irreversible aggregation and precipitation. The inventors hypothesize that the molecular interfacial layer between the aqueous and organic phases would protect the protein from prolonged contact with the organic phase. The inventors further hypothesize that PEG2KDSPE, being soluble in both layers, would remain at the interface, reducing the surface tension of the aqueous phase and allowing the protein to be electrospun into core-sheath filaments encapsulated by a sheath polymer. Previously, electrospinning of proteins was impossible due to immediate precipitation and aggregation upon contact between the two solutions, and the increased surface tension at the core-sheath interface. Experiment: This experiment evaluates whether including this excipient in the core can protect 20 mg / g of immunoglobulin protein from denaturation. Success metrics were: (a) visual assessment of the interface layer and evaluation for the presence of protein precipitation; and (b) assessment of protein integrity by size exclusion chromatography (SEC) of the % recovery relative to the expected value and the % HMW relative to the monomer percentage. All aqueous core solutions contained 20 mg / g of protein, which was first reconstituted with phosphate buffer at 50 mg / g. Data showed that increased concentrations of PEG 2k DSPE above 0.1% did not provide any additional protection at the core (aqueous) / sheath (organic) interface, and the surface tension increased to >41 mN / m with increasing PEG 2k DSPE, indicating that above 0.1% = CMC (critical micelle concentration).

[0119] Concentration: Solutions containing 0.25% and 0.4% PEG 2k DSPE, with and without methyl acetate, showed low protein recoveries, ranging from 60-79% at T=0 and even lower at T=2 hours. Solutions containing 0.1% PEG 2k DSPE, with and without methyl acetate, provided protection at the phase interface and achieved protein recoveries of 81-103%, remaining fairly consistent across the board for both T=0 and T=2 hours.

[0120] The interface between the aqueous and organic phases is addressed by the presence of PEG2K-DSPE in the aqueous phase and Span40 in the organic phase. The concentration of PEG-DSPE in the aqueous phase can be from 0.05% to 1%, and preferably from 0.1% to 0.4%. PEG-DSPE can be either PEG2K-DSPE or PEG5K-DSPE.

[0121] The (skin) polymer can 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, and mixtures and blends thereof, or a suitable organic solvent that dissolves the polymer and has a low diffusion coefficient into the aqueous phase. Depending on the polymer's molecular weight, the concentration of the polymer dissolved in the organic solvent is in the range of 5-50%, preferably 10-20%.

[0122] Skin polymers can also utilize the acetic acid component for dissolution to increase the polarity of the solution, provided the drug is compatible with acetic acid. For example, small molecule drugs are compatible with acetic acid in formulations, but large molecules such as proteins or DNA are not. Acetic acid can be a good choice for drugs that are stable at pH < 5.

[0123] The encapsulating (skin) polymer can be ethylene vinyl acetate (EVA) with a vinyl acetate content of 25-45%, of which 38-42% is preferred. The skin polymer can be a blend of EVA and other polymers to provide a resulting blend with the desired biomaterial properties. The skin polymer can also be polytrimethylene carbonate, or polytrimethylene carbonate-co-polycaprolactone, or polytrimethylene carbonate-co-DLL-lactide, which can be used as a single polymer or in suitable blends with other hydrophobic or hydrophilic polymers to impart strength, improve elasticity, and produce a porous or fracture-free adjacent encapsulating layer. Other polymers can be polyanhydrides, polyester-amides, polycaprolactone, polycaprolactone-TMC, polylactide, PLGA, and blends or copolymers thereof.

[0124] The implant contains a very high concentration of protein.

[0125] Proteins are contained within the middle layer of a "sandwich" insert. The types of proteins include immunoglobulins, globular proteins, recombinant proteins, fusion proteins, chimeric proteins, antibodies, structural proteins, growth factors, and enzymes. Stable formulations are those that maintain protein stability through their physical stability, chemical stability, and biological activity. Physical stability can be detected by size exclusion chromatography when precipitates (particles) or high-MW aggregates are formed. Chemical degradation can be size clipping, deamidation, methionine oxidation, disulfide scrambling, or other changes to the protein's primary structure. Both physical and chemical changes can reduce the biological potency of a protein.

[0126] One of the key challenges in developing dosage forms containing sustained-release proteins is their susceptibility to their microenvironment, which is determined by their concentration. Encapsulating high concentrations of proteins within a polymer matrix is ​​difficult. When dry implants are hydrated, protein concentrations in aqueous microenvironments (such as tears) are extremely high (>100 mg / mL), increasing the likelihood of both irreversible and reversible aggregation. The stability of proteins and other biopharmaceuticals is crucial for ensuring both safety and efficacy. The molecular structure of therapeutic proteins and key sites within these proteins are susceptible to oxidation, deamidation, hydrolysis, etc., and can lead to aggregation and fragmentation.

[0127] Therefore, the development of protein-based drug products begins with the work of stabilizing therapeutic proteins. The instability of proteins makes stabilization challenging; the folded state is only slightly more stable than the unfolded state, and any change in the protein environment can cause protein degradation, aggregation, and / or inactivation.

[0128] Temperature and pH are parameters that can affect protein stability, leading to thermal denaturation and aggregation, and chemical degradation such as side-chain oxidation, hydrolysis, and deamidation. For any specific protein, determining the pH window for maximum protein stability is crucial. The effect of pH on protein chemical stability can be altered in the presence of excipients. The effect of salts is largely dependent on the solution pH, affecting the charge state of ionizable groups in the protein. Depending on their type and concentration, metal ions can also stabilize or destabilize proteins, as their interactions with proteins are largely protein-dependent. Metal ions can also significantly affect protein stability without affecting most of its secondary structure. These target the Fenton pathway, specifically methionine, cysteine, histidine, tryptophan, tyrosine, proline, arginine, lysine, or threonine. Catalysis depends on the concentration of metal ions and can be promoted in the presence of reducing agents such as ascorbic acid salts or thiol (RSH) compounds.

[0129] Chelating agents such as ethylenediaminetetraacetic acid (EDTA) and citric acid can bind to proteins and destabilize them, or bind harmful metal ions and prevent oxidation. Mechanistically, protein concentration can also affect aggregation and, in some cases, chemical degradation. Higher concentrations increase the likelihood of undesirable self-association and consequently, higher viscosity.

[0130] Chemical degradation pathways include not only oxidation, hydrolysis, and deamidation, but also isomerization, succinimide formation and cleavage, non-disulfide crosslinking, and deglycosylation. Deamidation of asparagine and glutamine residues is the 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 sulfhydryl groups in methionine and cysteine ​​residues. Disulfide bond formation or sulfur-disulfide exchange can lead to protein aggregation or polymerization.

[0131] Soluble implant

[0132] Topical administration to the eye involves eye drops because these drops are rapidly cleared from the surface due to drainage from the nasolacrimal duct and overflow caused by tearing and blinking, resulting in low bioavailability (<10%). To minimize this effect, a slowly dissolving implant can effectively release polymeric excipients and / or drugs without overwhelming the ocular space.

[0133] This document discloses a variation of a water-soluble implant comprising both water-soluble and oil-based excipients, which is applied as a dry polymer implant to a tissue surface. Upon hydration, the implant absorbs water and transforms into a hydrogel, thereby conforming to the underlying tissue and optionally cross-linking to it. The transparent hydrogel forms an air-oil-water interface film on the hydrated tissue surface, thereby preventing water evaporation from the tissue surface. The hydrogel dissolves slowly based on its composition, and its dissolution time is modulated by the composition of the implant. The purpose of an oil-water implant is to form a “monolayer” interface on the tissue surface, in which oil floats on top, preventing evaporation.

[0134] In another variation, two different fibers are electrospun to form the implant structure. Upon contact with water, components in one fiber can react with components in the other fiber, thereby forming an in-situ hydrogel film on the tissue surface, which slowly dissolves / biodegrades in the biological environment.

[0135] Water-soluble implants can be used as bandages (without medication) to facilitate wound healing, or optionally contain a therapeutically effective pharmaceutical agent, a biopharmacologically active biological agent, or one or more polymers co-electrospun to crosslink with each other and with tissue. The dissolution time (from 1 hour to 2 weeks) can be varied by changing the type and concentration of the polymer contained in the electrospinning. Water-soluble implants can be applied to tissue surfaces. Types of tissues include the eye, nose, vagina, rectum, and oral cavity. Water-soluble implants can be applied as a bandage barrier to wounds or abrasions, optionally containing antimicrobial agents, analgesics, or drugs as wound-healing compounds. One example is a corneal bandage for wound healing of corneal abrasions. Its benefit is that, due to the membrane's transparency, there is no visual impairment caused by the device. Another example is an oral mucosal adhesive implant for covering oral ulcers while lubricating the oral cavity. One example is an oral implant containing ingredients to relieve dry mouth. Another example is a vaginal implant for relieving vaginal dryness. Another practical example of a soluble implant is a soluble implant that combines with clotting factors and can be filled into deep wounds to stop bleeding.

[0136] Characteristics of soluble implants

[0137] The dissolution time of the implant can be one hour, or it can be 12-24 hours, or even 1 to 14 days. Water-soluble implants can effectively deliver the ingredient to the ocular surface or any mucosal surface. The ingredient can include a drug or polymeric excipient that is released onto the ocular surface. The ingredient is a mixture of oil and aqueous polymer, which is emulsified and electrospun to form an air-oil-water interface. Implants are typically designed to be placed in the conjunctival fornix below the lower eyelid, but rapidly dissolving thin hydrogels can be placed on the sclera (the white part of the eye) for better absorption of the ingredient. For the treatment of ocular surface conditions, it is reasonable to place the implant on the corneal surface, such as... Figure 3 As shown in the image. The ocular implant can be placed into the conjunctival fornix (lower eyelid) for easy placement and rapid dissolution.

[0138] Variations of water-soluble, biodegradable implants are developed and formulated using biocompatible and tissue-soothing ingredients. The implants have a circular or curved shape to match the contours of the underlying tissue. The implants can be placed on the cornea to aid healing or to facilitate surgery. As dry devices, the implants are white and opaque, but can become transparent or translucent after hydration. The implants can be sterilized using low-temperature electron beam sterilization at a dose of 6–25 kGy and gamma radiation at a dose of 6–25 kGy.

[0139] The composition of the implant can be altered to change its dissolution time in the aqueous fluid of the eye. This can be achieved by incorporating water-soluble polymers of varying molecular weights. These polymers can be selected from, for example, polyethylene glycol, polyvinyl alcohol, hyaluronic acid, sodium alginate, xanthan gum, pectin, cellulose (carboxymethyl-, hydroxypropyl-, hydroxypropylmethyl-), guar gum, tamarind seed polysaccharide, 1->3β-glucan, poly(arginine), and polylysine to treat dry eyes, lack of lubrication, and itchy eyes caused by allergens or dryness. These implants are instantly soluble, dissolve in less than an hour, or are persistently soluble, with the dissolution time modulated by the implant composition. They can also relieve dryness in other tissues, such as dry mouth, as well as vaginal discomfort and dryness. The implants consist of one or more mucosal adhesive polymers and one or more biocompatible polymers. The biocompatible electrospun implant consists of two fibers with different compositions, one of which contains components that can optionally react with components in the other fiber. These fibers are co-spun simultaneously via electrospinning to form a water-soluble nanostructured implant. The implant's nanostructure is produced through electrospinning, electrospraying, or a combination of both. This nanostructure allows the implant to rapidly absorb water, optionally cross-link with each other and with underlying tissue. Through mild cross-linking, the implant does not immediately "ball" up into a gel mass but retains its shape. The dry device will dissolve in water and aqueous media, including biological media such as tears, nasal secretions, oral fluids, vaginal fluids, and rectal fluids.

[0140] Dry implant products can be prepared by electrospinning fibers. Two different fibers can be co-spun by an electrospinning process to form a single implant, wherein the first fiber contains a component that optionally reacts with a component contained in the other fiber. In one example, one fiber contains sodium hyaluronate, polysorbate 20 or polysorbate 80 with a MW in the range of 50,000-5,000,000 Daltons, and polyvinyl alcohol with a MW in the range of 20-130,000 Daltons, and optionally contains sodium chloride in the range of 0.1-3.5%, boric acid or its sodium salt in the range of 0.0001% to 5%, sodium alginate in the range of 0.01%-20%, and a component with a MW in the range of 0.01-60% of hyaluronic acid with a MW in the range of 100-100,000 Daltons. The implant contains polyethylene glycol in the Dalton range, dextran 70 in the range of 0.01-50%, calcium chloride in the range of 0.0001%-1%, polyvinylpyrrolidone in the range of 0.1-50%, and albumin in the range of 0.01-10%. The second fiber included in the implant contains polyvinyl alcohol with a MW of 20-130,000 Daltons, and optionally contains calcium chloride (0.0001-1%), sodium borate (0.0001-1%), and dextran 70. (1-50%), polyethylene glycol (MW 100-100,000 Daltons), PEG-multi-arm succinimide ester (arm number 4-16), polypropylene glycol (MW 400-4000), propylene glycol, type I collagen, glycerin, bioactive substances, sodium alginate in the range of 0.001-20%, sodium hyaluronate (MW in the range of 50,000-5,000,000 Daltons and in composition 0.01-1%), xanthan gum (0.01-10%), guar gum (0.01-10%), hydroxypropyl β-cyclodextrin, trehalose, mannitol, sorbitol, glycerin, tamarind seed polysaccharide or β1>6 glucan, wherein the first and second filaments are configured to react with each other and with biological tissue upon hydration, and the dry device transforms into a hydrogel upon contact with biological tissue and subsequent hydration, and after contact with biological tissue 0 Completely dissolves within 14 days.

[0141] The duration of dissolution can be varied by changing the concentration of the crosslinkable portion (e.g., alginate crosslinked with calcium chloride or other types of divalent salts such as magnesium chloride or zinc chloride). Another crosslinkable portion is boric acid crosslinked with polyvinyl alcohol or other polyhydroxy substances such as glycerol, sorbitol, or mannitol. In another embodiment of this concept, the concentration of polyoxyethylene polyarm amine can be varied with that of polyoxyethylene polyarm succinimide ester to form a loosely or tightly crosslinked hydrogel. It is not necessary for each fiber to be actually tightly crosslinked; optionally, each fiber can contain a slowly dissolving polymer, and if rapid dissolution of the dry device within 2 hours of placement on tissue is required, each fiber can be completely uncrosslinked.

[0142] For electrospinning solutions into fibers, the viscosity of PVA solutions at 25°C is in the range of 1000-5000 centipoise, and the viscosity of PEG solutions is in the range of 7000-15000 centipoise; the surface tension is 25-50 mN / cm, and the conductivity is 1-30 mS / cm.

[0143] In another example, the implant comprises hyaluronic acid salt and polyvinyl alcohol in a first fiber, and polyethylene glycol and alginate in another fiber. Boric acid and calcium chloride may be used in the fiber composition to crosslink the implant upon hydration.

[0144] The implant is applied as a dry device, rapidly absorbing water from the ocular surface or the soft tissue at its placement site. Upon hydration, the implant absorbs water from surrounding tissue to form a tissue-adhesive hydrogel.

[0145] An implant device made of electrospun fibers can optionally cross-react with components (including water) in the device to transform into a tissue-compliant hydrogel. Optionally, the implant may not contain components that react with each other. The two fibers may have the same or different components.

[0146] The hydrogel dissolves over time, with the dissolution time modulated by its composition. During dissolution, the hydrogel slowly releases its components over time. This slow release of the high-viscosity polymer components prevents the eye space from becoming overwhelmed, unlike eye drops that cause higher volumes of tears and tear flow.

[0147] Dry implants may contain tissue lubricating excipients, including but not limited to hyaluronic acid (0.05-15%, MW 80,000-3,000,000 Daltons), xanthan gum (0.05-15%), guar gum (0.05-15%), gelatin (0.05-15%), soluble collagen (0.1-17%), hydroxypropyl cellulose (HPC) (0.1-15%), hydroxypropyl methylcellulose (0.05-5%), lecithin (0.001-1%), dextran (0.05-15%), carboxymethyl cellulose (0.05-15%), glycerol (0.05-25%), tamarind seed polysaccharide (TSP) (0.05-15%), and 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, glycerin, propylene glycol, polypropylene glycol 400-4000, PEG 40-stearate, poloxamer, polyvinylpyrrolidone (PVP), tamarind seed polysaccharide (TSP), 1-3 Beta-glucan, pectin, polyacrylic acid, starch, gluconic acid chitosan, albumin, beta-glucan, gelatin, collagen, atelocollagen, lecithin, polyethylene glycol-modified lipids, distearylphosphatidylcholine, myristoylphosphatidylcholine, myristoyl / distearylphosphatidylglycerol, hydroxypropyl methylcellulose, hydroxypropyl cellulose, and mixtures thereof, at concentrations of 0.01-1%, 1-10%, and 10-15%.

[0148] The implant may contain water-insoluble excipients beneficial to the tissue, including oils (such as castor oil, peanut oil, olive oil, medium-chain triglycerides, corn oil, eugenol oil, sunflower oil, safflower oil, linseed oil, triglycerides, and diglycerides), menthol, camphor, ceramides, phospholipids, including derivatives and blends, and at a concentration of 0.01-30% in the final implant product. The benefit of incorporating water-soluble excipients into the water-soluble polymer solution is the creation of an air-oil-water interface to prevent water evaporation from the tissue surface. This is particularly useful in the treatment of dry eye. It is also useful in treating dryness in other tissues (such as vaginal dryness and dry mouth). The implant can be used to prevent wounds from drying out, as wound dryness promotes car formation.

[0149] The water-soluble implant may optionally contain a number of compounds, including but not limited to brimonidine, brinzolamide, timolol, prostaglandins, MMP modulators, antioxidants, olapatadine, antifungal agents, naphazoline hydrochloride, adrenaline, other vasoconstrictors, melatonin, lipoic acid, epitalon, n-acetylcysteine, glutathione, cysteine, superoxide dismutase, SOD mimics, aldehyde scavengers, vascular endothelial growth factor scavengers, immunoglobulins, RNA, DNA, peptide-DNA constructs, and coagulation. Blood factors, FGF, NGF, EDGF, anti-VEGF molecules, growth factors, cell-penetrating peptides, polyarginine, polylysine, albumin, heparin, n-acetylcysteine, amino acids, thymosin β, plant flavonoids, antibodies, ROS inhibitors, mast cell inhibitors, calcineurin inhibitors, PDGF inhibitors, analgesics, antiglaucoma agents, antihistamines, antiallergic drugs, vasoconstrictors, ultraviolet blocking compounds, polyphenols, wound healing agents, ROS quenchers, decongestants, lubricants, NO-releasing compounds, estrogens, estrogen mimics, antibacterial agents, antiviral agents, antifungal agents, antiemetics, analgesics, NSAIDs, analgesics, vitamins, oils, triglycerides, cholesterol, vitamin E, vitamin D, vitamin B12, DNA (including ministering DNA and plasmid DNA), RNA, and CRISPR-CAS9 constructs. Soluble implants can be used as artificial lubricants, slow-dissolving resources for eye lubrication, surgical aids, carriers for releasing molecules to relieve congestion or irritation, or emollients. Soluble implants can be used to release proteins, DNA, mRNA, RNAi, CRISPR constructs, or small molecules <1000 Daltons. Drugs can be incorporated into the implant, allowing for either pre-incorporation or no pre-incorporation.

[0150] The implant is prepared by electrospinning each fiber simultaneously (co-spinning) and collecting them on a collector, thereby forming a fiber mixture. Each fiber may be identical in composition, slightly different due to minor variations, or completely different. After hydration, the implant transforms into a tissue-compliant hydrogel with tensile strength and Young's modulus less than that of soft tissue.

[0151] Young's modulus is a very important factor to consider because biomaterials that are harder than the tissue being applied will cause irritation, and if they are not hard enough, they will not be able to maintain their structure.

[0152] For soft tissues and organs, Young's modulus ranges from 0.0001 MPa to 1 MPa. Specifically, the Young's modulus of the brain is 0.0001-0.001 MPa, that of the skin is 0.001-0.01 MPa, that of the spleen and pancreas is 0.0025-0.005 MPa, and that of glands and muscles is 0.008-0.017 MPa.

[0153] The Young's modulus of the liver is approximately 0.0105 MPa, and that of the human lateral nasal cartilage is 0.98 MPa. Furthermore, data on the ultimate tensile strength of soft tissues are also available. The conjunctival fornix has an ultimate tensile strength of 0.7 MPa and a Young's modulus of 3.9 MPa, while the amnion has an ultimate tensile strength of 1.7 MPa and a Young's modulus of 11.5 MPa.

[0154] like Figure 4 As shown, the Young's modulus of all dry inserts is less than 0.5 MPa.

[0155] The Young's modulus of the hydrated implant is in a similar range to that of the soft tissue, i.e., <0.5 MPa. Figure 5 This makes it possible to seamlessly integrate biomaterials with the underlying tissue.

[0156] The implant's internal microstructure is composed of two different types of fibers. The straight fibers are PEG-based, while the fibers with curved loops are alginate-based. Figure 6 Soluble implants can be sterilized by cryogenic electron beam or cryogenic gamma radiation.

[0157] This invention covers the following aspects:

[0158] Aspect 1: A dry drug delivery system comprising an intermediate layer enclosed within two outer layers, the intermediate layer and the two outer layers each comprising a mixture of electrospun fibers, the dry drug delivery system having a density of 0.1-1 g / cm³. 3The final total density within the range, thickness <1 mm, elastic modulus <4 MPa, and elastic strain change >100%, wherein each of the two outer layers comprises one or more polymers selected from the following: ethylene vinyl acetate, polytrimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) with a copolymer ratio of 90:10, PCL and P(TMC)-polyethylene oxide (PEG)-P(TMC) (wherein the molecular weight of the PEG segment is 1,000-20,000 Da and the molecular weight of the P(TMC) segment is 1,000-10,000 Da). Dal); and the intermediate layer is formed of core-shell type fibers, each of the core-shell type fibers comprising: a core containing a bioactive substance and one or more excipients that protect the bioactive substance from degradation, and an outer skin formed of a hydrophobic elastic polymer, the hydrophobic elastic polymer being composed of P(TMC) and PCL in a weight ratio of 90:10 to 30:70.

[0159] Aspect 2: The dry drug delivery system according to Aspect 1, wherein the vinyl acetate substitution degree of the ethylene vinyl acetate is 42%, the specific logarithmic viscosity of the P(TMC) is 0.3-1.2 dL / g, the specific logarithmic viscosity of the P(TMC)-co-PCL is 1.2-1.6 dL / g, and the copolymerization ratio is 90:10, and the specific logarithmic viscosity of the PCL is 1-1.5 dL / g, wherein the mixture of electrospun fibers in the two outer layers is annealed to form a continuous layer.

[0160] Aspect 3: The dry drug delivery system according to aspect 1 or aspect 2, wherein the mixture of electrospun fibers in the two outer layers is annealed to form a continuous layer.

[0161] Aspect 4: A dry drug delivery system according to any one of Aspects 1 to 3, wherein, based on the weight of the two outer layers, the 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).

[0162] Aspect 5: A dry drug delivery system according to any one of Aspects 1 to 4, wherein the system is biodegradable.

[0163] Aspect 6: A dry drug delivery system according to any one of Aspects 1 to 5, wherein the bioactive substance is a water-soluble drug selected from the following: small molecules with a molecular weight <1000 Dal, large molecules with a molecular weight >1000 Dal, proteins, amino acids, peptides or nucleic acids with a molecular weight of 10,000 Dal to 250,000 Dal.

[0164] Aspect 7: The dry drug delivery system according to aspect 5, wherein the nucleic acid is plasmid DNA, linear DNA, oligonucleotide, mRNA or siRNA.

[0165] Aspect 8: The dry drug delivery system according to aspect 1, wherein the bioactive substance is water-soluble and selected from the group consisting of: immunoglobulins, antigen-binding domains, antibacterial agents, antiviral agents, steroids, antiglaucoma agents, NSAIDs, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, antifungal agents, anti-inflammatory drugs, proteins, peptides, growth factors, enzymes, vitamins, hormones, polysaccharides, peptide-nucleic acids and nucleic acids or mixtures thereof.

[0166] Aspect 9: The dry drug delivery system according to aspect 1, wherein the bioactive substance is hydrophobic and selected from the following: antimicrobial agents, steroids, NSAIDs, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, antifungal agents, anti-inflammatory drugs, and antiglaucoma agents.

[0167] Aspect 10: The dry drug delivery system according to any one of Aspects 1 to 9, the dry drug delivery system further comprising a water-soluble adhesive fiber deposited on the system, the water-soluble adhesive fiber being selected from polyacrylic acid, PEG-polyarm amine, polyethyleneimine, polyamide amine, polylysine, polyarginine, glucuronide chitosan, and derivatives and blends thereof.

[0168] Aspect 11: A dry drug delivery system according to any one of Aspects 1 to 10, wherein during the preparation of the system by electrospinning, one or more excipients in the core of the core-shell type fiber interact with the bioactive substance as a water substitute, cooperate with the bioactive substance, or form an interface layer between the polymer-containing organic phase and the aqueous phase.

[0169] Aspect 12: The dry drug delivery system according to aspect 11, wherein the one or more excipients are selected from the group consisting of: hydroxypropyl β-cyclodextrin, γ-cyclodextrin, sulfobutyl β-cyclodextrin, PEG40 stearate, trehalose, sorbitol, sucrose, mannitol, poloxamer 407, poloxamer 188, polysorbate 80, polysorbate 20, dextran, polyvinyl alcohol, polyethylene oxide, polyvinylpyrrolidone, hydroxypropyl methylcellulose, human serum albumin, and soluble collagen, wherein the one or more excipients are present in the dry drug delivery system at a concentration of 1-75% w / w.

[0170] Aspect 13: The dry drug delivery system according to aspect 1, wherein the system has been sterilized by low-temperature electron beam radiation with a dose range of 6-25 kGy.

[0171] Aspect 14: The dry drug delivery system according to aspect 1, wherein the system is circular, oval, crescent-shaped, elliptical, annular, square, rectangular or tubular.

[0172] Aspect 15: The dry drug delivery system according to aspect 1, wherein the hydrophobic elastic polymer forming the outer sheath of the core-shell type fiber 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.

[0173] Aspect 16: A method for delivering a bioactive substance to the soft tissue of a subject, the method comprising: obtaining the dry drug delivery system of Aspect 1, and applying the system to the surface of the soft tissue before or after adding a hydration fluid to the soft tissue of the subject, the hydration fluid comprising an isotropic solution containing one or more of the following: sodium chloride, phosphate, citrate, albumin, magnesium salt, calcium salt, borate, boric acid, balanced salt solution, multi-branched PEG-amine, PVA, multi-branched polylysine, multi-branched polyarginine, guar gum, gellan gum, sodium alginate, xanthan gum, carboxymethyl cellulose, the hydration fluid having a pH of 6-7.8 and an osmolality of 270-340 mOsm / kg.

[0174] Aspect 17: A water-soluble drying 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 containing castor oil, buffer salts, amino acids, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinylpyrrolidone, dextran 70, albumin, PEG-polyarm amine, type I collagen, glycerol, or PEG40 stearate, and the second fiber containing polyvinyl alcohol (PVA) and sodium alginate, and optionally containing sodium borate, dextran 70, PEG, PEG-polyarm succinimide, polypropylene glycol, propylene glycol, type I collagen, bioactive substances, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or β-carotene. 1>6-glucan, wherein the first filament and the second filament are configured to react with each other and with biological tissue upon hydration, and the device transforms into a hydrogel upon contact with the biological tissue and subsequent hydration, and completely dissolves within 0 h to 14 days after contact with the biological tissue.

[0175] Aspect 18: The water-soluble drying apparatus according to Aspect 17, wherein the PVA has a molecular weight of 88,000-130,000 Da and a degree of hydrolysis of 88-100%, and the sodium hyaluronate has a molecular weight of 50,000-5,000,000 Da, and the sodium alginate has a molecular weight of 32,000-400,000 Da, the first fiber comprises 80-90% by weight of PEG, 0.5-5% by weight of sodium hyaluronate, 0.1-0.3% by weight of sodium chloride, and 0.001-2% by weight of polysorbate 20 or polysorbate 80, and the second fiber comprises 80-95% by weight of PVA.

[0176] Aspect 19: The water-soluble drying device according to aspect 17 or 18 further comprises a bioactive substance selected from the following: small molecules with a molecular weight <1000 Dal, macromolecules with a molecular weight >1000 Dal, proteins with a molecular weight of 10,000 Dal to 250,000 Dal, amino acids, peptides or nucleic acids of 2 to 100 amino acid residues.

[0177] Aspect 20: The water-soluble dry device according to aspect 19, wherein the bioactive substance is water-soluble, slightly water-soluble, or water-insoluble.

[0178] Aspect 21: The water-soluble dry device according to aspect 17, wherein, when placed on biological tissue, the device dissolves by simple dissolution in an aqueous fluid or by biodegradation.

[0179] Aspect 22: A water-soluble drying device according to any one of aspects 17 to 21, wherein the device is circular, oval, crescent-shaped, elliptical, annular, square, rectangular or tubular.

[0180] Aspect 23: The water-soluble dry device according to aspect 17, wherein the device has been sterilized by low-temperature electron beam radiation with a dose range of 6-25 kGy.

[0181] Aspect 24: The water-soluble dry device according to aspect 19, wherein the first fiber or the second fiber comprises the bioactive substance selected from immunoglobulins, antigen-binding domains, antibacterial agents, antiviral agents, steroids, antiglaucoma agents, NSAIDs, ultraviolet blockers, healing agents, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, analgesics, antifungals, anti-inflammatory drugs, proteins, peptides, growth factors, enzymes, vitamins, hormones, polysaccharides, peptide-nucleic acids, nucleic acids, or mixtures thereof.

[0182] Aspect 25: The water-soluble drying device according to aspect 21, wherein the biological tissue is an ocular mucosa, vaginal mucosa, rectal mucosa, nasal mucosa, or oral mucosa.

[0183] Aspect 26: A method for delivering a bioactive substance to the soft tissue of a subject, the method comprising: obtaining the water-soluble dry device of aspect 17, and applying the system to the surface of the soft tissue of the subject, wherein the device transforms into a hydrogel upon hydration upon contact with the surface of the soft tissue.

[0184] Without further elaboration, it is believed that those skilled in the art can utilize the invention to its fullest extent based on the above description. Therefore, the following specific embodiments should be interpreted as illustrative only and not as limiting the remainder of this disclosure in any way. All publications (including patent documents) cited herein are incorporated herein by reference in their entirety.

[0185] Example

[0186] Example 1: Soluble alginate / polyvinyl alcohol insert

[0187] This is an example of a cross-linked soluble implant formed from two solutions with different compositions. Solutions 1 and 2 were electrospun into two different fibers.

[0188] Table 1: Polymer Solution #1

[0189]

[0190] Table 2: Polymer Solution #2

[0191]

[0192] Table 3: Composition of the Dry Device Insertion Components

[0193]

[0194] In dissolution studies, the implant crosslinked upon hydration (within 5 minutes of soaking in the buffer) in phosphate-buffered saline at pH 7.4 and 37ºC, transforming into a tissue-fit hydrogel. The hydrogel dissolved within 21–36 hours. In terms of composition, the implant was designed to allow alginate (via Ca2+) to crosslink. 2+ Both cross-linked PVA (through borate cross-linking) and PVA (through borate cross-linking) form a more durable interpenetrating network. Dissolution time can be shortened by reducing the concentrations of borate and calcium chloride. In the absence of calcium chloride and borate, the PVA / alginate / PEG35K implant dissolved in vitro in less than 8 hours. In another example, an implant consisting solely of PEG35K was electrospun, and the implant body dissolved within 30 minutes.

[0195] Example 2: Soluble Hyaluronic Acid / Alginate / PVA Implant

[0196] In this embodiment, the implant that forms the hydrogel is prepared by electrospinning, which involves co-spinning two fibers from two different solutions simultaneously.

[0197] The implant 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).

[0198] Solution 1 contains 35,000 MW polyethylene glycol (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 has a pH of 5.5–6.0, a conductivity of 1.154 mS / cm, a surface tension of 31.59 mN / m, and a viscosity of 9264.4 cP.

[0199] Solution 2 contains PVA88 (8.97%), sodium alginate LVP (0.5%), polysorbate 20 (0.20%), acetic acid (0.146%), sodium chloride (0.40%), sodium hyaluronate (0.5%), and endotoxin-free water. Solution 2 has a pH of 3.9–4.2, a conductivity of 5.23 mS / cm, a surface tension of 35.12 mN / m, and a viscosity of 1918.8 cP.

[0200] The final dry implant composition (by weight) is 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.

[0201] The implant formed a hydrogel within 1 minute and became completely transparent within 30 minutes. The implant existed as a thin hydrogel at 4 hours and dissolved within 4–6 hours. The use of calcium chloride in solution 1 caused mild cross-linking of sodium alginate in the implant, while boric acid in solution 1 caused cross-linking of PVA in solution 2. Adjusting the calcium chloride and boric acid in the implant composition reduced the cross-linking density of the implant and thus shortened the dissolution time. The control implant formed without the use of calcium chloride and boric acid dissolved rapidly in less than 3 hours.

[0202] Example 3: Soluble collagen / polyvinyl alcohol-based in-situ mucosal adhesive implant (using hydration solution)

[0203] In this embodiment of the soluble implant, the implant comprises type I collagen fibers, PVA, and polysorbate 20. Upon placement on tissue, the mesh implant can be in situ crosslinked to form a tissue-adhesive hydrogel by application of a borate-phosphate saline hydration solution. The implant is designed to dissolve upon placement on tissue. The hydration solution components induce the crosslinking reaction. The implant is prepared using an electrospinning process from a single solution containing type I collagen (4-8%), 88% hydrolyzed PVA 130K (50-90%), and polysorbate 20 (0.1-2%) (see Table 4 below).

[0204] Table 4. Solution composition of SB-NANOM-X-11-34

[0205]

[0206] To produce collagen solutions of any composition, acetic acid must be used to create a pH <3.5. The solution used had a pH of 3.3. Conductivity: 1875 µS / cm; Viscosity: 571.6 cP. After electrospinning, the PVA / collagen mesh inserts are in situ cross-linked on the tissue using borate in the hydration solution. The hydration solution was maintained at a pH of 7.4–7.5 using a borate-phosphate buffer.

[0207] In dry implants, the concentration of small molecule drugs can range from 0.1% to 10%, with PVA content ranging from 50% to 95% and collagen content ranging from 3% to 8%. The polyvinyl alcohol (PVA) MW can range from 23,000 Da to 200,000 Da, and the degree of hydrolysis is 88% to 100%. 100% hydrolyzed PVA contains no vinyl acetate and has very low solubility in water. PVA was purchased from Sigma-Aldrich; type I collagen was purchased from Biomatrix, Nutragen Cat# 5010 (6 mg / mL, in 1 N HCl, pH 3.5) and Fibricol, Cat# 5133 (10 mg / mL, in 1 N HCl, pH 3.5).

[0208] The water-soluble compounds used in these studies had water solubility >100 g / mL and MW <500 Da. The compounds were ionizable at pH >4.5. The hydrogel formulation was subjected to monolithic electrospinning. The hydrogel was formulated to consist of collagen (which is biocompatible and integrates structural integrity) and PVA (which facilitates electrospinning by increasing the viscosity of the solution).

[0209] Dry web formulations are the final formulations of webs produced after the electrospinning process.

[0210] Table 5. Dry Reticulated Formulation of SB-NANOM-X-11-34

[0211]

[0212] Electrospinning: During the electrospinning process, the solution formulation contaminates the needle tip at the conventional voltage range (<18 kV). As a result, only a small amount of fiber is formed, but for the majority, the formulation is electrosprayed. When the voltage is increased to 19 kV, the Taylor cone stabilizes, with only slight gelation at the tip. However, fibers are then produced.

[0213] Table 6. Electrospinning parameters of SB-NANOM-X-11-34

[0214]

[0215] Use of hydration solution: An isotonic borate-phosphate aqueous solution is used as the hydration solution to crosslink the components of the hydrogel implant and fix the drug delivery system to the underlying tissue. Crosslinking of the implant is achieved using 20 mM, 25 mM, or 40 mM borate buffer. The hydration solution is first applied to the tissue, followed by the application of an 8 mm implant. Without the use of borate-phosphate buffer, the implant is hydrated but not fixed to the mucosal surface. Therefore, the borate-phosphate hydration solution is an essential design element for collagen / PVA-based rapid dissolution systems. Preparation of hydration solution: The wetting solution contains borate, while the phosphate component contributes to pH stability. Borate ions crosslink with the PVA component in the implant formulation ( Figure 7 Cross-linking mechanisms are crucial for the bioadhesion and structural integrity of implants. Phosphate buffer was added to the wetting solution to stabilize the solution's pH, as pH 6.8 is outside the pKa range of 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 to achieve the final osmotic pressure.

[0216] Table 7. 25 mM borate-3 mM phosphate buffer preparation

[0217]

[0218] Table 8. Properties of 25 mM borate-3.3 mM phosphate buffer

[0219]

[0220] Rapid hydration: The collagen / PVA-based implant is initially white and dry. After application of the implant, a 40 μL droplet of hydration fluid can be applied, inducing the hydration process. In another embodiment of this concept, the hydration fluid can be applied before or after the implant is placed onto the tissue. Upon placement, the implant begins to hydrate, rapidly absorbing water and transforming into a hydrogel. The hydrogel appears to "sink" into the tissue, molding into the crevices. This exemplary implant has a thickness of 60 micrometers. After overnight incubation in a buffer solution, the hydrogel is no longer visible on the tissue. At pH 7.8, there is no significant difference in crosslinking rates between borate buffer solutions of different strengths. The thickness of the dry device can range from 60 micrometers to 1 mm, with a preferred thickness of 60-1000 micrometers and a most preferred thickness of 100-450 micrometers. For some applications, the thickness of the implant can be up to 2000 micrometers, although these applications are for non-ophthalmic indications. When collagen / PVA-containing implants were placed on 30 µL of a 25 mM borate / 3 mM phosphate saline solution in in vitro experiments, the initially smooth, opaque, white implants (8 mm in diameter) absorbed water and transformed into spherical hydrogels. The time taken for the implants to form a complete hydrogel structure was recorded and is shown in Table 1. Each sample shrank upon initial contact with the hydration solution and formed a hydrogel, a result of the cross-linking mechanism between PVA, glycerol, collagen, and borate ions. The cross-linking reaction between PVA and borate ions is well known, but the interpenetrating network between PVA, glycerol, and collagen forms a hydrogel that can form in situ on the tissue surface. Implants containing only PVA collapsed into a “slime” upon cross-linking with borate, lacking an integral structure. PVA / collagen / glycerol implants have a smooth texture and can deliver drugs to the site and dissolve into non-toxic components. Therefore, the inclusion of collagen is an essential design element. Hydrogel formation time: The hydrogel formation time depends on the thickness of the implant. Sample KR-NANOM-X-10-61-02 is the thinnest, with a thickness of 0.006 mm. After the implant was placed on a droplet of hydrated solution, most of the implant dissolved in just 43 seconds, leaving a small piece to crosslink and form a hydrogel. The remaining samples with larger thicknesses took more than 1 minute to form hydrogels and did not dissolve rapidly. See Table 9 below.

[0221] Table 9. Time required for the network to completely transform into a hydrogel

[0222]

[0223] Shape Retention: Each sample shrinks and floats on the surface of the borate buffer, maintaining its round shape. The inclusion of collagen is an important design feature for maintaining the implant's shape. Collagen is insoluble in water at pH > 7, thus forming a water-swelling hydrogel. Solubility Study: The solubility of the implant was determined upon contact with a 25 mM borate / 3 mM phosphate aqueous solution. Figure 8 (Cross-linking kinetics and dissolution of collagen-containing hydrogel formulations). The newly formed hydrogel was then diluted with 20 mM phosphate-buffered saline solution, partially mimicking the environment around the conjunctiva of the fornix. Samples were placed in a monitored 37°C incubator at 0.5 h, 1 h, 2 h, 4 h, and 6 h. The physical state of the solution and hydrogel placement was observed and reported, and the solution was measured for drug quantification. ~100% of the drug was recovered from the hydrogel.

[0224] The 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, partially mimicking the environment around the conjunctiva. Samples were placed in a monitored 37°C incubator for the following time points: 0.5 h, 1 h, 2 h, 4 h, and 6 h. After the addition of 20 mM phosphate buffer, the implant remained a hydrogel and floated on the surface of the solution, exhibiting a lower density compared to water. Vials were placed in the incubator and removed at their respective time points.

[0225] After 64 hours of incubation, a thin, translucent membrane was left in 20 mM phosphate buffer, which was used to simulate the environment in ocular tissue.

[0226] Figure 8 The dissolution of the collagen-containing hydrogel implant is shown. Dissolution kinetics are faster in tears containing lipids, enzymes, and mucins. It was also concluded that implant thickness is related to the time in 'hydrogelation time'. See Table 9 above.

[0227] Example 4: In vitro and in vivo studies of a monolithic electrospun implant containing PVA / collagen

[0228] The implant was formulated by adding acetic acid to make the solution more compatible with the collagen component (which was supplied as a 0.01M HCl solution) and to increase conductivity. The solution was significantly clearer and more homogeneous. The model drug was water-soluble and had a molecular weight <500 Da.

[0229] Table 10. Solution formulations with batch number SB-NANOM-X-11-75

[0230]

[0231] Table 11. Dry Reticulated Formulations with Batch Number SB-NANOM-X-11-75

[0232]

[0233] Electrospinning parameters: The insert is integrally electrospun, which implies the use of a syringe and needle containing a single solution to spin a mesh. During the electrospinning process, bead-like fibers and droplets were observed to emerge from the needle tip at the start of the process. The initial humidity (67%) decreased to 35%. Due to the humidity change, droplets from the needle tip were trapped. The mesh appeared uniform, and the drug compound exhibited an average encapsulation recovery value of 103% (n=3, STD: 7%). This demonstrates that the mesh is uniform in terms of API loading.

[0234] Table 12. Electrospinning parameters of SB-NANOM-X-11-75

[0235]

[0236] In vitro studies:

[0237] In vitro studies were conducted to determine the bioadhesive and hydrogelation properties of SB-NANOM-X-11-75 on bovine eyelids. Samples were placed on the eyelids and a borate wetting solution was applied (see [link to study]). Figure 9 (Application of hydrogel implant) to activate the cross-linking mechanism between the PVA component of the implant and the borate component in the wetting solution.

[0238] like Figure 10 As shown, a hydrogel forms approximately 20 seconds after the wetting solution is applied. The tissue and implant are gently pushed together with tweezers. The implant moves with the tissue, indicating adequate bioadhesion.

[0239] For example Figure 10 As shown, after 90 seconds of wetting, the implant remained adhered to the tissue and retained its round, translucent properties. Forty minutes after application, the implant was a clear hydrogel, still intact.

[0240] Example 5: In vivo study of SB-NANOM-X-11-75

[0241] After application to the rabbit eye, the implant transformed into a hydrogel and did not cause discomfort in the animals for up to 6 hours, with no excessive blinking, rubbing, or discharge. However, the implant was not visible, meaning it may have become embedded in or fused with the tissue. The implant was not visible at the 1-hour or 6-hour time point. See also Figure 11 The implant is applied to the conjunctival fornix; Figure 12 Application of wetting solution; Figure 13: The hydrogel insert after wetting; and Figure 14 (The implant is inserted into the tissue one hour after it is placed in either eye).

[0242] In in vivo studies, batch number SB-NANOM-X-11-75 appeared to lack structural integrity, as it dissolved within the first hour after application to the conjunctival fornix. This is likely due to insufficient cross-linking caused by uneven or ineffective application of the borate solution. The hydrogel cannot achieve full structural integrity when the PVA-borate mechanism of action does not occur uniformly throughout the implant and tissue. Therefore, further optimization of the hydrogel formulation and electrospinning process is needed.

[0243] Example 6: In vivo and in vitro studies of co-spun soluble implants

[0244] SB-NANOM-X-11-75 exhibited insufficient cross-linking, and the implant dissolved within the first hour of in vivo administration. This is likely a result of a lack of matrix structural integrity. Based on results from in vivo studies, the hydrogel implant was reformulated to prolong its dissolution time.

[0245] Example 7: Soluble, in-situ crosslinkable PEG35K / PVA100 / PVA88 / alginate implant

[0246] This embodiment is an in-situ formed hydrogel implant, which forms upon hydration by tissue fluid. This implant does not require a hydration fluid. After placement, the implant slowly hydrates with tissue fluid, inducing an interlocking, cross-linking mechanism that anchors the implant to the underlying mucosa. The hydrogel nature of the implant contributes to comfort and tactile sensation. The hydrogel implant is transparent, disappears into the tissue, and is invisible to the naked eye. The dissolution rate can be varied depending on the specific concentrations of alginate and polyvinyl alcohol.

[0247] The implant device is formed by co-electrospinning of two solutions. One solution contains a treatment-relevant concentration (1-35%) of the drug, PEG35K (10-60%), calcium chloride (up to 0.5%), and boric acid (0.01-5%), while the other solution contains 0.01-0.5% polysorbate 20, 0.40-0.50% sodium alginate, and 7-25% PVA100 / PVA88. Glycerin may be included as a surface wetting agent, and other buffer salts, such as phosphates, acetates, histidines, or citrates, may be used.

[0248] In one specific example, the nanofiber implant is formed by co-spinning a first solution and a second solution. The first solution contains 55% PEG 35kDa, 0.14% boric acid, 0.05% calcium chloride, and 0.01% polysorbate 20 in water for injection or distilled water. The second solution contains 2.6% PVA100, 7.7% PVA88, 0.44% sodium alginate, 0.1% polysorbate 20, 0.5% acetic acid, and 10% pharmaceutical ingredient. The first solution has a viscosity of 23,184 cP, a conductivity of 56.8 µS / cm, a pH of 6.2, and a surface tension of 45 mN / m. The second solution has a viscosity of 2270 cP, a conductivity of 565 µS / cm, a pH of 4.3, and a surface tension >45 mN / m.

[0249] The dry implant consists of 15.8% PEG 35 kDa, 0.028% boric acid, 0.01% calcium chloride, 0.003% polysorbate 20, 10.3% PVA 100, 31% PVA 88, 1.78% sodium alginate, 0.42% polysorbate 20, 0.7% glycerol, and 40% model compound.

[0250] In another example, the implant is formed by co-spinning a first solution and a second solution, the first solution containing 35-55% PEG 35kDa (Merck lot number 57742792 924, cat#8.18892.1000) in water for injection (QS) or distilled water; 0.14% boric acid (boric acid: Spectrum lot number: 2GE0219, cat# B0120, NF, EP, BP, JP grade); 0.15% calcium chloride (calcium chloride (dihydrate): Millipore Sigma lot number A0300782 629, cat#1.02382, ACS, EP grade); 0.1% polysorbate 20 (Tween 20; Croda, super refined, SR40800), and the second solution containing PVA 100 (6%) and PVA 88. The first solution contained 3.2% sodium alginate (0.46%), polysorbate 20 (0.01%), acetic acid (0.5%), and a drug (6%). The viscosity of the first solution was 24,000 centipoise (CP), measured at 25°C and a shear rate of 40° (Anton Paar rheometer MC92, axis: Part# 18163, diameter 25 mm, angle: 1°); conductivity: 279 µS / cm (Mettler Toledo model#); pH: 4.63; and surface tension >50 mN / m. The second solution contained 24,000 cP, conductivity: 225 µS / cm, pH: 4.63; and surface tension >45 mN / m.

[0251] The dry implant consists of 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, and 35% pharmaceutical substances.

[0252] Typically, solution 1 has a viscosity of 15,000-30,000 cP at 25°C, a conductivity of 50-350 µS / cm, and a pH of 4-7. Solution 2 has a viscosity of 1100-2500 cP, a conductivity of 500-1000 µS / cm, and a pH of 4-8.

[0253] In in vivo trials in rabbit eyes, the hydrogel implant was deemed comfortable and dissolved slowly over time.

[0254] Example 8: Co-spun soluble implant

[0255] Two phases (solutions) are required for co-spinning the solution used to process SB-NANOM-X-11-91. Sodium alginate is added to the co-spinning formulation, along with calcium chloride. However, after several trials, sodium alginate did not remain in the solution with collagen due to pH incompatibility. Sodium alginate (pH 5-9 in water) and collagen (pH 4 in HCl / water) do not form a stable solution because neither molecule requires the pH necessary to remain stable in an aqueous environment. Sodium alginate is a carbohydrate that cross-links via reaction with calcium ions, while the hydrogel formation mechanism is due to PVA forming cross-links via reaction with borate ions. Collagen was removed because it did not contribute to promoting cross-linking or bioadhesion. The implant prototype does not require wetting.

[0256] Table 13. Solution #1 SB-NANOM-X-11-90-1

[0257]

[0258] Table 14. Solution #2 SB-NANOM-X-11-90-2

[0259]

[0260] Table 15. Dry Reticulated Fabric SB-NANOM-X-11-91

[0261]

[0262] Electrospinning of both solutions is possible. The viscosity of the PEG phase (solution #1, batch number -90-1; see Table 13 above) should be reduced to improve flow and alleviate pressure on the pump during future spinning. Encapsulation analysis of the implant (mean value, n=3, 108% drug recovery) demonstrates the homogeneity of the network.

[0263] Table 16. Electrospinning parameters for preparing SB-NANOM-X-11-91

[0264]

[0265] Table 17. ECU / Pin Location

[0266]

[0267] Table 18. Physical characteristics of SB-NANOM-X-11-91 dry insertion component

[0268]

[0269] The thickness of the implants in batch-91 ranges from 130 to 144 µm, and the density ranges from 0.216 to 0.254 mg / mm². 3 The radius of the insert is 3 mm.

[0270] Example 9: In vitro study of SB-NANOM-11-91

[0271] Freshly excised calf conjunctival tissue was obtained and used for evaluation of implant batch-91 within one day of excision. Hydrogel formation was observed after placement onto the moist conjunctival tissue. Following initial placement, the implant hydrated slowly, inducing an interlocking, cross-linking mechanism that anchors the implant to the underlying mucosa (see [link to relevant documentation]). Figure 15 As in previous in vitro studies, the hydrogel was gently pushed into the tissue using clean tweezers, and it adhered to the tissue and maintained its round shape. After 24 hours, the implant was not visible to the naked eye (see [link to previous study]). Figure 16 ).

[0272] Example 10: More durable formulation: Soluble, in-situ crosslinkable PEG35K / PVA100 / PVA88 / alginate implant

[0273] More durable hydrogels can be formulated by including glycerol in a solution containing 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 whether the hydrolysis state of PVA affected the dissolution of the implant. Glycerol was also included in the formulation as a surface wetting agent. The prototype was co-spun, as in SB-NANOM-X-11-91, meaning that the crosslinking mechanism does not require a wetting solution to occur.

[0274] Tables 19 and 20 below show two solutions co-spun using the same process used to prepare 11-91. However, the addition of sodium alginate and CaCl2 components to enhance crosslinking in 15-14 reduces the dissolution rate. Process control characteristics are shown in Table 21, and the dry insert composition is shown in Table 22.

[0275] Table 19. Solution #1 SB-NANOM-X-15-14-01

[0276]

[0277] Table 20. Solution #2 SB-NANOM-X-15-14-02

[0278]

[0279] Table 21. Process Control Characteristics

[0280]

[0281] Table 22. Dry Reticulated Formulations

[0282]

[0283] The electrospinning parameters are shown in Table 23 below. Note that the viscosity of the PEG phase can be slightly lower to establish improved flow, which will reduce the pressure on the pump, but overall, the plume is narrower and the process is stable.

[0284] Table 23. Electrospinning parameters of SB-NANOM-NANOM-X15-14

[0285]

[0286] Example 11: In vivo study of SB-NANOM-X-15-14

[0287] The structural integrity of the implant device from batches 15-14 was assessed to determine whether alterations to the concentration of specific components in the formulation (i.e., the concentration of PVA (polyvinyl alcohol) and the sodium alginate content) would affect hydrogel dissolution upon implantation into the rabbit eye fornix. General observations of congestion, comfort, and irritation were made at specified time points T=1 hour and 6 hours.

[0288] Results of in vivo studies in New Zealand white rabbits Figure 17 and 18 As shown in the figure. The implant exhibited: (a) rapid hydration to form a transparent hydrogel, (b) formation of a device with adhesive properties to the conjunctiva, and (c) no grooming or squinting behavior in the animal after placement, demonstrating that the hydrogel device is comfortable. At both 1 hour and 6 hours, the implant was completely dissolved, and no residue was observed.

[0289] During in vivo studies, prototype #2 (15-14) was applied to each eye of one rabbit (see [link]). Figure 19 The 6-hour time point was the endpoint of the study. (See application...) Figure 20 Following administration of the SB-NANOM-X-15-14 implant, a very small amount of eye discharge was observed on the ocular surface near the corner of the eye close to the nose, but this was considered minor. After sedation (6 hours later), a hydrogel was observed; this hydrogel was slightly opaque, viscous, and spherical (see [link to product description]). Figure 21 (Examination of the ocular surface, T=6 hours). The results showed that 15-14 (which was formulated using two types of PVA and whose main components were completely hydrolyzed) did not achieve complete dissolution in vivo for up to 6 hours.

[0290] Example 12: Longer-lasting formulation: The role of glycerol added to the formulation of a soluble, in-situ crosslinkable PEG35K / PVA100 / PVA88 / alginate implant.

[0291] The objectives of the network #SB-NANOM-X-15-22 are: (i) to increase the PVA content to increase the density of the PVA phase, which can help form a more uniform final network; (ii) to adjust the PVA blend to PVA100 / PVA88 (25:75). The 65:35 blend, consisting of more fully hydrolyzed PVA, makes the implant dissolve too slowly. Reducing the amount of fully hydrolyzed PVA can increase the dissolution time of the implant (up to 6 hours; the current target is 3 hours); and (iii) to reduce the CaCl2 and sodium alginate levels to the concentrations shown in 11-91 to obtain less crosslinking, thereby enabling the implant to dissolve within the 3-hour time point.

[0292] Table 24. Solution #1 (PEG phase) of SB-NANOM-X-15-22

[0293]

[0294] Table 25. Solution #2 (PVA phase) of SB-NANOM-X-15-22

[0295]

[0296] Table 26. Dry Reticulated Formulation of SB-NANOM-X-15-22

[0297]

[0298] Stability studies were conducted on samples (mesh batch number SB-NANOM-X-15-22) that had been exposed to electron beam sterilization. This experiment evaluated the effects of 15 and 25 kGy electron beam exposures on drug-containing soluble implants.

[0299] All samples were analyzed after one month of stability at 5°C, 25°C / 60% RH (relative humidity), and 40°C / 75% RH. Samples exposed to a 15 kGy electron beam were adequately sterilized without compromising the properties of the implant.

[0300] Table 27. Sterilization by electron beam

[0301]

[0302] As shown in Table 27 above, the implant properties were maintained after irradiation with a 15 kGy electron beam. In another test, the implant device was sterilized by 9 kGy gamma radiation under dry ice. The implant properties were maintained.

[0303] The concept of a soluble implant can be used to deliver medications or healing ingredients to the ocular surface to treat conditions on the ocular surface. These conditions can include corneal conditions such as keratoconus, dry eye, blepharitis, Sjogren's syndrome, corneal abrasions and tears, corneal lesions, bacterial and fungal infections, allergic conjunctivitis, and post-cataract surgery inflammation prevention. Eye conditions can include blepharitis, meibomian gland cysts, conjunctivitis, contact lens-related problems, corneal abrasions, corneal dystrophy, corneal edema, corneal erosion, corneal ulcers, dacryocystitis, ectropion, endophthalmitis, entropion, episcleritis, ocular tumors, foreign bodies, fungal keratitis, glaucoma, Graves' ophthalmology, hypotony, keratitis, migraine, mucormycosis, optic nerve retinitis, oculomotor palsy, optic nerve-related problems, orbital cellulitis, actinic keratitis, scleritis, sinusitis, stye, surgery, ocular burns, trauma, uveitis, UV damage, glaucoma, and corneal vascularization. Dissolvable implants can also be used as a treatment for Sjögren's syndrome of mucous tissues from other routes (such as the nose, rectum, mouth, and vagina). The soluble implant can be used with or without medication, including small molecules, proteins, peptides, hormones, adenovirus-based drugs, RNAi, and CRISPR-CAS9. The implant can be used to treat brain diseases, particularly for oncology applications.

[0304] Example 13: Preparation of a soluble implant using water-soluble and insoluble (oil) components

[0305] These inserts are designed to lubricate the eye surface, thereby creating an air-oil-water interface and preventing water evaporation.

[0306] Table 28. Composition and properties of solution IG-NANOM-DSW-02-75A

[0307]

[0308] Table 29. Composition and properties of solution IG-NANOM-DSW-02-77A

[0309]

[0310] Example 14: Preparation of a cross-linkable and soluble dry implant containing 39-amino acid peptides

[0311] The soluble implant described in this embodiment is composed of two different fibers: one containing a multifunctional PEG8-arm-NHS and the other containing a PEG8-arm-NH2. The implant is a dry device that absorbs physiological moisture from the soft tissue surface. Crosslinking occurs at pH > 7, which causes the fibers and tissue to crosslink with each other, resulting in a slowly dissolving implant. The solutions are prepared according to the compositions in Table 30 below. In this embodiment, two solutions, each containing a multifunctional polyethylene oxide (PEG8-arm-amine or PEG8-arm-NHS), are prepared. (NHS = succinimidyl ester; GAS = glutaramide succinimidyl ester) Each solution contains ethanol, THF, polyvinyl alcohol, and gellan gum. The pH of each solution is estimated to be 6-7. The PVA solution (10%, in water) and the gellan gum solution (1.5%, in water) are mixed at room temperature with stirring (PVA / gellan gum 80% / 20% w / w) for 20 minutes. As part of the preparation of solution 1, API (exenatide) is added to a portion of the PVA / gellan gum solution. The mixture is vortexed for 5 min. PEG-amine dissolved in ethanol / THF is added to the above solution, and the mixture is vortexed for 5 min to complete the preparation of solution 2. PEG-NHS and Tyloxapol dissolved in ethanol / THF are added to another portion of the PVA / gellan gum solution. The mixture is vortexed for 5 min to form solution 2. Ethanol and THF are added to increase the dielectric constant (conductivity) of the solution. The pH of each solution is estimated to be 6-7. The solution is loaded into a 3 mL syringe and connected to an injection pump that will deliver the solution at the same rate. The distance between the syringe and the tip is approximately 5 cm. The electrospinning equipment is manufactured by Tong Li Tech Co, Ltd; model TL-01. This equipment is set up for electrospinning by providing high-voltage power by connecting wires to the syringe tip. A connecting roller or plate is placed in the chamber and grounded. The parameters are as follows: voltage 9.5kV, roller RPM 30, pump flow rate 0.3 mL / h, and distance between tip and end roller / plate 10 cm. Start the syringe pump and turn on the voltage for electrospinning. Spin the solution into a continuous insert. Dry the insert overnight in a vacuum oven at room temperature.

[0312] Table 30. Composition of electrospinning solutions 1 and 2

[0313]

[0314] Wetting with hydration buffer: The following hydration solutions are used for dry-device implants with ring fibers, wherein the outer layer is hydrophilic and reacts with each other, while the inner layer is hydrophobic and contains a drug. The purpose of the hydration solution is to: (A) wet the tissue surface to establish maximum contact between the implant and the tissue, and (b) induce a cross-linking reaction between the polymerized amine (NH2) incorporated into the implant and the polymerized succinimide ester (NHS) and / or proteins on the tissue surface. The high pH of the hydration solution induces the cross-linking reaction. Examples of hydration solutions are prepared with the following compositions, some of which are shown in the table below. The hydration solution contains formulations of sodium borate, sodium chloride, and gellan gum. The force required to “peel” the implant from the tissue after binding (1–30 minutes) is the peeling force (N).

[0315] Table 31. Composition of hydrated solutions

[0316]

[0317] Other hydration solution variations may contain 0.1-0.5% xanthan gum, 0.1-0.5% guar gum or 0.1-0.5% hydroxypropyl guar gum, 0.1-0.5% balanced salts, 0.1-0.5% polycarbophil, and 0.1-0.5% sodium hyaluronate.

[0318] Table 32. Insert compositions for hydration and peel bond strength

[0319]

[0320] Solutions 1 and 2, as shown in Table 32 above, were co-spun onto the rotating collector of an electrospinning apparatus. Co-electrospinning of solutions 1 (in a 3 mL syringe) and 2 (in a 5 mL syringe) was performed under the following conditions: voltage: 9 kV; flow rate of both syringes: 0.2 mL / h; collector roller speed: 40 RPM. YL-NanoM-1-32 was used as the co-spun insert, in which reactive PEGs (see Table 32, solutions 1 and 2) were each in separate solutions to prevent premature reaction. The two solutions were spun into monolithic fibers, each containing one of the reactive PEG components. Peptides have previously been shown not to react with PEG-amines, and were therefore included in solution #1. In the presence of the hydrated solution, an electrophilic-nucleophilic reaction between the PEGs was induced to form a tissue-bound hydrogel.

[0321] Quantitative analysis of the mechanical adhesive strength of the implant to the tissue

[0322] The mechanical adhesion of the hydrated implant to conjunctival tissue was evaluated using a Shimadzu mechanical tensile testing instrument. The tissue was fresh bovine conjunctiva. The implant was YL-NanoM-1-32. The dry implant was applied to the fresh bovine conjunctival tissue and then hydrated using 40–50 μL of hydration solution. The mechanical testing instrument was attached to one end of the implant. The hydrated implant / tissue was fixed to a horizontal substrate. The force required to separate the implant from the tissue was recorded as N (stress) / mm (strain) (peel test).

[0323] like Figure 22 As shown, complete bonding occurs approximately 30 minutes after the implant is applied to the conjunctival tissue and then moistened with hydration buffer (YL-NanoM-pep-1-59-6; see Table 33 below). The peel force (N) at 30 minutes is 0.028N–0.05N. The peel force of dry implants (without hydration solution) and implants hydrating immediately after placement on the tissue is 0.008N. This allows for removal of the implant in case of improper application. The electrospun implant transforms into a clear hydrogel. Peel strength tests were conducted using a peel test model to compare the peel strength of the various hydration solutions listed in Table 33.

[0324] Table 33. Peeling force (N) of fresh bovine conjunctival implants as a function of different hydration solutions.

[0325]

[0326] The hydrated solution YL-NanoM-pep-1-59-6 contains both sodium borate and gellan gum, and exhibits the highest peeling force from tissue implants. Figure 22 (Right side). Hydration experiments were conducted using YL-NanoM-1-32 (see Table 32).

[0327] Kinetics of peel strength of bovine conjunctiva using inserts with gellan gum / borate pH 7.8 buffer

[0328] The kinetics of peel strength between the implant and bovine conjunctiva were measured. The implant was cut into 1×2 cm pieces (YL-NanoM-pep-1-79-3; see composition in Table 34 below) and applied to fresh bovine conjunctival tissue. Fifty μL of hydration buffer was added to the implant, and peel strength was measured over time. The hydration solution consisted of 30 mM sodium borate buffer and 0.25% gellan gel. The results are shown in Table 35 below.

[0329] Table 34. Composition of YL-NanoM-pep-1-79-3 (Desiccant Placebo Insert)

[0330]

[0331] Table 35. Kinetics of the binding of hydrophilic sheath implants to bovine conjunctiva

[0332]

[0333] Longer bonding time and higher peel strength were observed between the implant and the tissue. It was also found that placing the dry implant first, followed by the hydration solution, provided superior bonding compared to applying the hydration solution first to the tissue.

[0334] Kinetics of peel strength between the implant and bovine conjunctiva using polymer / borate buffer-based buffer

[0335] In vitro bioadhesion tests were performed using fresh bovine conjunctiva (using a Shimadzu Instron tester) using the electrospun implant YL-NanoM-pep-1-79-3 (see Table 34). An 8 mm implant patch was placed onto the tissue and allowed to adhere without pressure. After application, 40 μL of hydration buffer was applied to the patch to hydrate it. All tests were performed with a 5-minute adhesion timeframe.

[0336] Table 36. Composition of hydration buffer solution

[0337]

[0338] Table 37. Peel strength of electrospun implants from bovine conjunctival tissue using gellan gum / borate, sodium hyaluronate / borate, and polycarboxylic acid / borate (pH 7.8).

[0339]

[0340] The data in Table 37 above show that implants hydrated with sodium hyaluronate / borate and polycarboxylic acid / borate at pH 7.8 have higher peel strength compared to implants hydrated with gellan gum / borate buffer at pH 7.8.

[0341] Other polymers that may be used in the hydrated solution include xanthan gum / borate, polyacrylate / borate, gellan gum / borate, guar gum / borate, PolyQuad-1 / borate, polylysine / borate, phosphatidylcholine / borate, chitosan / borate, polyarginine / borate, PAMAM / borate, PEI / borate, trilysine / borate, and tetralysine / borate. Mannitol, sucrose, and trehalose may be used as optional additives. Tyloxacillin / borate, polysorbate 20 / 80, poloxamer, octaxynol, and triton-X100 may be used as optional surfactants, added at concentrations ranging from 0.01% to 1%, depending on permissible FDA limits and the required concentrations to achieve the necessary energy supply. The concentration of the polymer may be from 0.05% to 0.5%. Polymer concentrations below 0.05% are inefficient, while concentrations above 0.5% tend to form viscous gel solutions, which are unsuitable for dispensing into eye drop bottles. Sodium chloride is added to equilibrate the solution to an osmotic pressure of 280-320 mOsm / kg.

[0342] In another exemplary implant, the mutually reactive polymers PEG10K-8-arm-NH2 and PEG10K-8-arm-NHS are incorporated with PVA into a single solution. This is typically difficult to achieve because the components of the solution are mutually reactive. However, reactivity can be minimized by mixing the cold solution and maintaining a relatively low pH. The solution is adjusted to pH 5–5.5 using hydrochloric acid (see Table 38). This prevents premature crosslinking of the mutually reactive polymers.

[0343] Table 38. Composition of the solution for combined PEG implants

[0344]

[0345] The effect of pH

[0346] The composition of the combined solution can be within a pH range of 4-5.5. At pH < 4, degradation of PEG10K-8-arm-GAS (NHS) occurs due to the hydrolysis of ester bonds. At pH > 5.5, premature reactions occur between PEG components, therefore, the solution cannot be electrospun. Within a pH range of 4-5.5, the hydrolyzable bonds in PEG10K-GAS are stable, making this an ideal pH range.

[0347] Effect of PEG concentration: The composition of the combined solution in the electrospinning solution can be increased or decreased within the range of 0.5% to 10%. Higher concentrations of activated PEG result in stronger tissue "anchoring" and a longer time for hydrogel dissolution after application of the hydrated solution. In dry device inserts, the concentration of activated PEG is 5-50%.

[0348] Adding 200 μL of ethanol to 1 g of solution facilitates the electrospinning of fibers. The dry insert of YL-NanoM-pep-2-21 consists of 84.85% PVA 67K, 7.58% PEG8-arm-NH2 (10K), and 7.58% PEG8-arm-NHS. Spinning conditions are: needle tip size: 20; flow rate: 0.5 mL / h; voltage: 9.5 kV.

[0349] The reaction between borate and polyvinyl alcohol:

[0350] The reaction between borate and polyvinyl alcohol (PVA) is known. However, an in-situ reaction on tissue allows the PVA-borate interpenetrating network to interlock with the tissue, thereby anchoring the implant to the tissue surface. Therefore, the coexistence of PVA and PEG in the implant composition enables two different mechanisms of tissue bonding. Peel strength measurements were performed on the hydrated implant YL-NanoM-pep-2-21. The implant was flexible in the dry state, with a tensile strength of 1.2–6 N / mm. The implant thickness was approximately ~200 mm. A 1 × 2 cm portion of the implant was crosslinked using a borate buffer hydration solution. The hydration solution was 0.25% gellan gum / 30 mM borate at pH 7.8. The implant was hydrated directly on fresh bovine conjunctiva. After a 30-minute crosslinking time, a peel test model of the implant was performed using a mechanical testing instrument. Based on multiple measurements, the peel strength of hydrated YL-NanoM-pep-21 is 1.4-3 N.

[0351] This enhanced adhesion of the implant to tissue is extraordinary and addresses the residence challenge of local sustained-release drug delivery systems. Therefore, the retention of the implant on the tissue surface for the duration of release is a key quality characteristic for the success of a drug delivery system.

[0352] In another experiment, PEG8-arm-NH2 and PEG8-arm-NHS were tested in acidified water at higher concentrations up to 15% w / v to evaluate their adhesion strength to tissue. The adhesion strength to tissue increased to >6N due to a higher cross-linking density of the network formed between PVA, the PEG, and endogenous proteins on the tissue.

[0353] In another experiment, the molecular weight of PEG in the implant was changed to 20K 8-arm PEG. The adhesion strength of the hydrated implant to the tissue was measured by a peel test and confirmed to be >1N, which is the minimum standard for good adhesion.

[0354] In another experiment, the pH of the solution was further reduced to 4.5 and 5. The adhesive strength of the hydrated implant was measured by a peel test and determined to be greater than >3N.

[0355] To further optimize the peel strength, the composition of the hydration solution was evaluated. High adhesive-peel strength was achieved by increasing the borate concentration to 50 mM. At 50 mM, the adhesive-peel strength was >3 N.

[0356] Integrity of the encapsulated peptides in the implant:

[0357] In this experiment, we determined whether peptides could remain intact during the electrospinning process. 10–20 mg of drug-loaded inserts were sonicated in 1 mL of aqueous buffer to dissolve all components. Ethanol was added to precipitate the polymer, and the resulting slurry was centrifuged. The supernatant was collected for HPLC analysis (RP C18 column; see HPLC methods in Table 39 below), and the peptide content was analyzed.

[0358] Table 39. HPLC methods for peptides

[0359]

[0360] Measurement of encapsulated peptides in the implant: 10-20 mg of the drug-containing implant product was sonicated in 1 mL of aqueous buffer to dissolve all components. The drug content was determined by HPLC analysis (RP C18 column), and the peptide content was also analyzed. Therefore, the implant contains 1.6-2.8% peptides, or 16-28 µg / mg.

[0361] Table 40. Encapsulation of peptides (APIs) in implants

[0362]

[0363] The compositions of YL-NanoM-1-32 and -34 are given in Table 32 above. Solutions 1 and 2 were simultaneously electrospun to form a drug delivery system on the same substrate.

[0364] Peptide integrity: Degradation products of peptides extracted from the implant were also analyzed and characterized as % oxidation. The peptides maintained their integrity throughout the encapsulation process. (See also...) Figure 23 ).

[0365] In vivo evaluation of soluble hydration implants

[0366] The in vivo behavior of the hydrated peptide implant was evaluated. Figure 24 A hydrated implant adhered to rabbit eye tissue. The tested implant was compositionally identical to YL-NanoM-pep-02-13, except for the absence of any proteins (see Table 41 below). The implant appeared translucent, highly hydrated, and adhered well to the ocular mucosa. No hyperemia or conjunctival edema was observed at T=0 or at the end of 5 days. The implant's hydration percentage was approximately 65%–75%.

[0367] In vivo stimulation assessment of peptide-soluble implants

[0368] An in vivo study of the peptide-containing implant was conducted in six normotensive New England white rabbits. A sterile implant (placebo implant) was inserted into the conjunctival fornix of each eye. Baseline hyperemia in each eye was assessed before and after implant placement. Rabbits were assessed twice daily for five days. General stimulation assessments were performed. After five days, no hyperemia or erythema was observed in any of the treated eyes. Therefore, the implant was considered biocompatible.

[0369] Irritation testing of soluble implants using the EpiOcular Surrogate Model

[0370] Irritation tests were performed using the Epiocular Corneal Construct Model (Cyptorex, LLC) on implants with a composition similar to YL-NanoM-pep-1-32 (see Table 32 above) but lacking any proteins, and compared with benzalkonium chloride as a control. These coated hydrated placebo implants were classified as non-irritating, while the benzalkonium chloride solution was considered highly irritating. The coating on the implants was a PEG-based hydrogel, while the core matrix was PLGA-based.

[0371] In vivo residence study of soluble implants

[0372] In an in vivo residency study in New Zealand white rabbits, placebo implants with 10% placebo concentration of 8-arm activated PEG (PEG8K-NH2+PEG8K-NHS) were tested in slightly acidified water (pH 4.5), n=6, for 14 days. Optimized borate buffer (pH 7.8) was used to hydrate the implants on tissue. Hydrated implants remained in place for 7–10 days (the study ended at day 14).

[0373] In another equivalent study, in rabbits (n=3, or 6 eyes), implants containing highly 8-arm functionalized activated PEG (PEG10K 8-arm-NH2 + PEG10K-8-arm-NHS) and IgG maintained adhesion to the conjunctival surface for up to 21 days. In another study, 16-arm functionalized activated PEG implants could remain on mucosal and conjunctival tissue for more than 2 weeks.

[0374] In vitro release (%) of peptides from soluble implants

[0375] At 37ºC, 20–40 mg of sample was placed in a 1 mL float-a-lyzer tube, suspended in 40 g buffer, and swirled. Samples were collected at time points during the study period and analyzed by HPLC. Figure 25 The release of exenatide (a highly water-soluble 39-amino acid peptide) from electrospun inserts of different compositions was demonstrated. The insert YL-NanoM-pep-1-64 (which has the same composition as YL-NanoM-pep-02-13 except that exenatide was used instead of BSA and IgG) showed a much slower rate of diffusion from the dialysis tube compared to the control peptide in aqueous solution (buffer solution, pH 7.4).

[0376] Encapsulation of other peptides in an electrospun implant has been achieved. The implant was prepared using an electrospinning process and encapsulated in a blend of PLGA (RG504 + poly(TMC-LL), 50:50). The solvent was methyl acetate, and the polymer concentration was 20%. The peptide was thymosin β-4, a protein encoded by the TMSB4X gene in humans. The human protein consists of 43 amino acids (SDKPDMAEIEKFDKSKLKKTETQEKNPLPSKETIEQEKQAGES (SEQ ID NO: 1)) and has a molecular weight of 4921 g / mol. After preparation, the implant was dried under vacuum to remove residual solvent. Encapsulation (mg / G): 20 mg of the drug-loaded implant product was dissolved in 1 mL of solvent to dissolve all components. 1 mL of aqueous solution was added to precipitate the polymer. 1 mL of the slurry was centrifuged, and the supernatant was used for HPLC analysis (RP C18 column) to analyze peptide content. Peptide encapsulation is expressed as µg peptide / mg product. Peptide integrity: The drug product was dissolved in an organic solvent, and degradation products were analyzed and characterized by % oxidation.

[0377] In vitro release and 1-hour burst rate (%): At 37ºC, (20–40 mg) of sample was placed in a 1 mL float analysis tube, suspended in 40 g buffer, and swirled. Samples were collected at time points throughout the study period and analyzed by HPLC. Figure 26 The diagram illustrates different release modes of the three prototypes. As shown, small water-soluble peptides can be released from the implant in a continuous manner. Other peptides, both hydrophobic and hydrophilic, can be incorporated into the implant, including melatonin, estrogen peptides, collagen peptides, epitaron, thymosin B4, and thymosin β4 fragments.

[0378] Implant Irritation Test of Crosslinked Soluble Implants: An in vivo study using implant YL-NanoM-pep-02-13 was conducted in normotensive New England white rabbits (Toxikon Corp, Bedford, MA). The sterile implant was inserted into the conjunctival fornix of each eye in the rabbits. Baseline hyperemia in each eye was assessed before and after implant placement. Rabbits were evaluated for 2–3 hours. For the placebo implant at maximum thickness, no hyperemia, animal discomfort, or irritation was observed in the eyes of normotensive rabbits during the duration of the study.

[0379] Encapsulation of IgG and BSA in PEG-GAS / PEG-NHS cross-linked implants:

[0380] Bovine IgG (Sigma Aldrich) (a protein with a molecular weight of ~120,000 g / mol) was encapsulated in a soluble insert. Analysis of the IgG showed that the protein was completely encapsulated and that the IgG was completely released over time.

[0381] In another example, an implant containing both IgG (MW 120,000 g / mol) and bovine serum albumin (BSA; MW 62,000 g / mol) was prepared. The composition is shown in Table 41 below.

[0382] Table 41. Composition of implants containing two different proteins

[0383]

[0384] The in vitro release results are shown in Table 42 below. Both IgG and BSA were released in a sustained manner.

[0385] Table 42. In vitro release of two bioactive proteins from soluble PEGGAS / PEGNHS implants

[0386]

[0387] Example 15: Incorporating intact protein into the implant at a high concentration

[0388] These experiments demonstrate that certain excipient mixtures stabilize proteins at very high concentrations (>100 mg / g). Proteins are unstable molecules and readily aggregate at very high concentrations. Therefore, their microenvironment is crucial in their stabilization strategies. Proteins are also prone to aggregation and eventual degradation due to contact with organic solvents, especially in the presence of both aqueous media and organic solvents.

[0389] To achieve high protein loading and maintain release in electrospun implants, a protein solution was loaded as the "core" aqueous solution and a polymer as the "sheath" organic solution in a core-skin type electrospinning apparatus. After electrospinning, the protein was incorporated into the core channels, which were then surrounded by a hydrophobic sheath. The compatibility of different solvents and polymers used to prepare the prototype electrospun mesh implant with IgG protein was evaluated. Excipients stabilizing high protein concentrations were determined in the following experiments.

[0390] 100 mg / g IgG was prepared in 50 mM phosphate buffer (pH 7.4); 10 mg / g PVP, PVA, and PEG were prepared in 50 mM phosphate buffer (pH 7.4). Organic / water / IgG solutions were prepared by mixing IgG (10 mg / g, in buffer) with 40% v / v ethanol, IPA, DMSO, and methyl ethyl ketone (MEK) in a 1:1 ratio. Samples at T=0 were analyzed by size exclusion chromatography using UV-Vis detection at 220 nm. The remaining formulations were incubated at 40°C and sampled on day 4 by repeating the previous steps (T=4 days). Table 43 summarizes the compatibility solutions.

[0391] Table 43. Compatibility Studies

[0392]

[0393] As shown in Table 43, PEG (polyethylene glycol), PVA (polyvinyl alcohol), PVP (polyvinylpyrrolidone), PAA (polyacrylic acid), borate buffer, phosphate buffer, Triton X-100, ethyl acetate, DMSO, and poly(LLA-TMC) are compatible with IgG. After lyophilization and reconstitution with water to 100-130 mg / g, these excipients contribute to the stability of both BSA and IgG, enabling the formulation of proteins at high concentrations. In another study, 0.3-1 mg / g of IgG was incubated (or held at room temperature) with TFA (trifluoroacetic acid) and polyacrylic acid. Analysis by SEC HPLC showed ~100% recovery.

[0394] In another study, 100 mg / g BSA was incubated with 8.5 mg / g and 18 mg / g arginine. For solutions containing higher arginine concentrations, ~100% recovery was observed. Other excipients, HPβCD (80–200 mg / g) and PEG2KDSPE (1–10 mg / g), exhibited 100% protein recovery when incubated with BSA (100 mg / g).

[0395] In another study, ~100% recovery was achieved for 160–240 mg / g IgG in phosphate-buffered saline (PFS) using 2–10 mg / g HPβCD, 18 mg / g arginine, and 160 mg / g HPβCD. IgG (250 mg / g in PFS) incubated overnight in several separate solutions (containing 20–40 mg / g HPβCD in PFS, 40 mg / g HPβCD in PFS and ethanol, or 40 mg / g mannitol in PFS) showed recoveries ranging from 98–100%. In another study, the effects of dextran, PEG40 stearate, Tween 20, and sorbitol were investigated (see Table 44 below). These excipient combinations stabilized high concentrations of IgG (100 mg / g) with recoveries of 98–102%.

[0396] Table 44. Excipient Combinations of Dextran, PEG40 Stearate, and Tween 20

[0397]

[0398] Polymers in organic solvent combinations (such as n-methylpyrrolidone, Span40, THF, anisole, dichloromethane, methyl acetate, ethyl acetate, butyl acetate) incubated with IgG in buffer solution do not degrade proteins, thus enabling the incorporation of very high concentrations of protein (100-150 mg / g) into electrospun inserts. IgG and albumin at concentrations of 50-150 mg / g dissolved in phosphate buffer (pH 6-7.5) and containing the following components are long-term stable: polysorbate 20 (0.1-1%), polysorbate 60 (0.1-1%), polysorbate 80 (0.1-1%), polyvinyl alcohol (0.5-100 mg / g), polyvinylpyrrolidone, 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 cyclodextrin (10-200 mg / g), and sulfobutyl β-cyclodextrin (10-200 mg / g). 5-20 mg / g), hydroxypropyl methylcellulose (10-20 mg / g), hydroxypropyl cellulose (10-20 mg / g), hydroxyethyl cellulose, tamarind seed polysaccharide (5-200 mg / g), β-glucan (5-200 mg / g), xanthan gum (5-20 mg / g), hydroxypropyl guar gum (0.1-10 mg / g), alginate (0.5-10 mg / g), hyaluronic acid (0.1-2%).

[0399] Hydrophobic polymers compatible with IgG and capable of being spun into nanofiber implants (which contain proteins) are poly(trimethylene carbonate) (1-20%), poly(caprolactone) (1-20%), poly(lactic acid-co-glycolic acid) (1-25%), poly(sebacic anhydride) (1-20%), polyanhydride (1-25%), polyorthoester (1-25%), polyester-amide (10-25%), poly(organosilicon) (1-30%), polyurethane, polyester-polyurethane, polyanhydride-polyurethane, ethylene vinyl acetate polymers having a vinyl acetate content of 20-50%, and blends thereof.

[0400] Protein stability under alkaline conditions (pH 7-8) is maintained by reducing the number of negatively charged amino acids (Asp, Glu) and increasing the number of neutral hydrophilic amino acids (His, Asn, Gln, and Arg). It has been shown that amino acids in the range of 0.1-10% (arginine, glycine, proline, glutamine, histidine, cysteine), such as low molecular weight excipients like sucrose and sorbitol, can improve the stability of bovine serum albumin and IgG. Polyamino acids (polylysine, polyglutamate) at concentrations of 0.1-5% in aqueous buffers (pH 6-7.5) can stabilize proteins (IgG, growth factors).

[0401] Other proteins stabilized in the above combinations include nerve growth factor, fibroblast growth factor, epidermal growth factor, recombinant antibodies, therapeutic antibodies, antibody-like scaffold proteins, glycoengineered proteins, and immunoglobulins. Among therapeutic proteins, anti-inflammatory drugs, Fc fusion proteins, anticoagulants, blood factors, bone morphogenetic proteins, engineered protein scaffolds, enzymes, growth factors, hormones, interferon, interleukins, thrombin, fibrin, fibrinogen, thrombolytics, and thrombolytic agents can be dissolved in high concentrations and stabilized by the combination of these excipients in the aforementioned amounts.

[0402] Other proteins that can be incorporated into implants include interleukin receptor antagonists, immunoglobulins, VEGF inhibitors, aldose reductase inhibitors, antihypertensive drugs, antioxidants, growth factor agonists and antagonists, vitrectomy agents, adenosine receptor antagonists, adenosine deaminase inhibitors, glycosylation antagonists, anti-aging peptides, topoisomerase inhibitors, antimetabolites, 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 σ ligands, endothelin ETA / receptor antagonists, antihyperglycemic agents, antiglaucoma agents, antichromatin-modifying enzymes, insulin, glucagon-like peptides, immunosuppressants, tissue repair agents, essential fatty acids, and nucleic acids (such as plasmid DNA, linear DNA, stranded DNA, peptide nucleic acids, antisense oligonucleotides, mRNA, siRNA, and RNAi).

[0403] The following peptides can be incorporated into electrospun implant products: exenatide, cyclosporine, bacitracin, vancomycin, daptomycin, voclosporin, glutathione, oxytocin, insulin, glucagon, bivalirudin, buserelin, corticotropin, cosyntropin, enfuvirtide, eptifibatide, glatiramer, epitarone, melatonin, melatonin, adiponectin peptide, and their derivatives.

[0404] Other peptides that can be incorporated into electrospun implants include defensin peptides, cell-penetrating peptides, anticoagulant peptides, milk-derived peptides, antimicrobial peptides, collagen peptides, GLP-1 peptides, PPAR inhibitory peptides, adiponectin-targeting peptides, WNT-targeting peptides, EGFR-targeting peptides, hormone peptides, tumor-related peptides, and TAT peptides. Other peptides already incorporated into delivery systems include pineal peptides, including but not limited to epitalon and melatonin.

[0405] Example 16: Preparation and characterization of IgG-containing implants (hydrophilic core-hydrophobic skin)

[0406] In the absence of a top and bottom layer of sandwich structure, or in the presence of a top and bottom layer of sandwich structure, proteins can be contained within a core-skin structure.

[0407] The goal is to produce defect-free implants with high protein loading and integrity. The IgG-containing implant is constructed from two types of electrospun fibers, one of which contains a core-sheath type fiber containing the protein. For the core-sheath type fiber (similar to a coaxial cable), the "sheath" is hydrophobic and polymeric, acting as an encapsulation barrier to slow the diffusion of water-soluble proteins (enclosed in the core), thus enabling sustained protein release. In the absence of the core-sheath structure, protein release from the monolithic fiber (monolithic = single-fiber microstructure) is faster (100% release within 1 hour at 37ºC).

[0408] The core solution (aqueous) (before electrospinning) contains a hydrophilic polymer that may be selected from the following list: polyvinyl alcohol (1-30%), polyvinylpyrrolidone (PVP) (1-20%), polyethylene glycol 2000-10,000-distearate (PEG-DSPE; 1-50%), PEG40-stearate (1-20%), dextran (1-3%), and sorbitol (10-25%). The sheath polymer solution (organic phase) used is: PDLLA-TMC, poly(TMC) or poly(TMC)-co-poly(caprolactone) and poly(caprolactone) (inherently 20% w / w solubility) dissolved in methyl acetate / DMSO at a ratio of 90:10, in a 50:50 ratio, or PCL+PTMC-PCL (15-30%, in methyl acetate / DMSO). Span 40 is used as a surfactant in this formulation to create an organic-water interface, thereby minimizing protein aggregation. Suitable molecules for creating an organic-water interface are dipalmitoylphosphatidylcholine (DPPC), distearate phosphatidylcholine (DSPE), phosphatidic acid, Tween-like molecules, Span-like molecules, ceramides, and polyethylene glycol-modified lipid molecules.

[0409] During electrospinning, the core solution feed rate is 0.1–5 ml / h, while the sheath solution feed rate can be 0.1–10 ml / h. Advantageously, the relative humidity varies between 20% RH and 40% RH, and the temperature varies between 22°C and 30°C. The distance between the collector and the spinneret is 10 and 18 cm, and the voltage difference between the spinneret and the collector is 10–18 kV.

[0410] Table 45 below shows core-sheath compositions for coaxial protein-containing products, wherein the protein is located in a hydrophilic core and the polymer is located in a hydrophobic sheath. Each of the solution compositions for the hydrophilic core and hydrophobic sheath is provided. The hydrophobic polymer content in the dry device is 50-95% w / w. PTMC = polytrimethylene carbonate; PCL = polycaprolactone.

[0411] Table 45. Core-sheath compositions for IgG pharmaceutical products (IgG is located in a hydrophilic core, and the polymer serves as the encapsulation membrane sheath).

[0412]

[0413]

[0414]

[0415]

[0416] Example 17: Encapsulating intact proteins in implant products

[0417] like Figure 28 As shown, the standard (1 mg / g IgG standard), the pre-encapsulation solution containing 11% IgG (core), and IgG extracted from the implant product exhibited similar levels of protein monomers (~73-77%), as well as high molecular weight fractions (14-21%) and low molecular weight fractions (0.9-2.5%). The data demonstrate the effectiveness of encapsulating high concentrations of protein in the implant product.

[0418] Example 18: Release of intact proteins from implanted drug products

[0419] Figure 29 The figure illustrates the release of intact protein (represented as % monomer) and other substances (represented as HMW1, HMW2, and LMW). Comparison of 1 mg / g standard and released protein at each time point up to 192 hours (8 days) shows that the released protein at each time point was intact and consistent with the standard.

[0420] The data also demonstrate that highly water-soluble proteins (IgG) are continuously released from implants with a thickness of less than 250 micrometers. Figure 29 The IVRT data for the implant shown is for a coaxially spun implant with a sheath polymer (PCL and PCL-TMC (10:90) blended in a 50:50 ratio) dissolved in methyl acetate, and a core solution of 0.1% SPAN 40 and IgG (in PEG40-stearate) and 0.1% polysorbate 20 in 20 mM phosphate buffer (pH 7) (concentration 11%). The data demonstrate that it is feasible to incorporate high concentrations of protein into thin implants without resulting in high aggregation fractions (HMW1, 2). As shown in the figure, no low molecular weight degrading proteins were observed (the peak at residence time ~13 min is attributed to polysorbate 20). In contrast, the web spun without polysorbate 20 or PEG40 stearate in the core protein solution exhibited very low protein recovery (<20%) after encapsulation. In another instance, implants containing sorbitol in a protein-containing core solution also exhibited excellent protein recovery (>90%). Figure 30 It also demonstrated the sustained release of proteins from the implant product.

[0421] Example 19: Release of intact proteins from implanted drug products

[0422] Figure 29 and 30The data demonstrates the continuous release of proteins from the implants. Both implants were coaxially spun using an electrospinning device.

[0423] The protein is contained in a core solution that also contains hydroxypropyl β-cyclodextrin, which may bind to the protein to provide protection against aggregation. The sheath polymer is polycaprolactone (PCL12) co-dissolved in methyl acetate at a concentration of 20% and blended with poly(trimethylene carbonate-polylactide) (90:10) in a 50:50 ratio, containing 0.1% SPAN 40. In contrast, the protein aqueous solution is emulsified in an organic polymer solution (composed of the aforementioned sheath solution), resulting in a burst release of the protein from the implant.

[0424] Figure 30 Sustained release from poly(trimethylene carbonate)-co-poly(lactide) fibers was demonstrated. However, the implants did not exhibit dimensional stability, were very stiff, and therefore were not biocompatible with tissues. Release patterns on multiple implants were not reproducible, which was more due to subtle differences in the manufacturing process compared to release patterns obtained using sandwich implants containing poly(TMC)-co-poly(caprolactone) (90:10).

[0425] Figure 31 High protein recovery rates were demonstrated in multiple batches using coaxial electrospinning with an aqueous protein solution as the core and an organic polymer solution as the sheath. The data demonstrate the practicality and reproducibility of the protein-containing core-sheath implant product. In contrast, batches without PEG40-stearate, PEG2K-DSPE, 0.1% polysorbate 20, or sorbitol (0.1-11%), or 1-50% HPβCD consistently exhibited recovery rates <20%. This demonstrates that the amphiphilic surfactant creates a protective interface between the aqueous protein solution (core solution) and the organic polymer solution (sheath), thereby preventing denaturation events that lead to aggregation. Figure 31 The relative recoveries for different batches are shown as percentages. Batches 12-995A, 12-995B, and 12-665C all contain the aforementioned excipients, but at concentrations <0.2%. This results in relatively low percentage recoveries.

[0426] Example 20: Sterilization method for protein implants

[0427] A key requirement for protein products intended for ocular or in vivo administration is sterility. Protein solutions for injection and infusion are typically sterile filtered unless the high concentration of protein or excipients makes the solution too viscous for sterile filtration. IgG-containing implants are prepared using the components shown in Table 47 and are electron beam sterilized.

[0428] Table 47. Protein-containing implant products used for electron beam sterilization tests

[0429]

[0430] Electron beam sterilization of products containing protein implants was performed while the products were cooled on dry ice at electron beam intensities of 15 kGy, 25 kGy, and 40 kGy (Ebeam Services, Inc.). Following sterilization, the protein content of the mesh implants was analyzed by size exclusion chromatography (SEC). The results are shown in Table 50 below.

[0431] Table 48. Protein recovery as a function of electron beam sterilization

[0432]

[0433] The implants shown in Table 47 contain high concentrations of IgG (9%) in the core solution, as well as Tween 20 and sorbitol as protein protective excipients. The components in the sheath solution also show to be protein-compatible. Data in Table 48 show that electron beam exposure at 25 and 40 kGy is detrimental to the encapsulated protein (% protein recovery ~50%-68%), while exposure at 15 kGy shows 96-100% protein recovery. Other sterilization methods proven to have protein retention capabilities are NO2 sterilization (Noxilizer, Inc.) (96% protein recovery), X-rays, and gamma radiation at doses of 1-15 kGy. Sterilization methods without retention capabilities include gamma sterilization at doses >15 kGy (<10% recovery) and ethylene oxide (ETO). With the use of high electron beam exposure, there is a significant increase in low molecular weight fractions (left side, right side). Figure 32 ).

[0434] Example 21: Selection of a size-stable polymer for constructing drug-containing hydrophobic fibers

[0435] The polymer, including the hydrophobic fibers of the implant, needs to maintain its size and shape after placement on ocular tissue. Implants that do not exhibit dimensional stability may stretch and separate from the tissue when placed on conjunctival tissue, causing mucosal adhesion failure and discomfort.

[0436] Experiment #1: The implant was constructed from the following materials: (a) polycaprolactone (90:10) blended with poly(trimethylene carbonate):polycaprolactone (Sigma-Aldrich) in a 50:50 ratio, and polycaprolactone (PC12); (b) polylactide-co-glycolic acid, 5004 (Purac); (c) polydioxanone (PDO); and (d) polycaprolactone, 80K (PCL 80K). Results: Upon hydration, the PLGA implant shrank in volume and became harder in texture and feel. In contrast, the PCL / PTMC blend (polycaprolactone 80K blended with 90:10 polytrimethylene carbonate:polycaprolactone) (Sigma Aldrich), PDO (polydioxanone, Sigma Aldrich), and PCL 80K maintained dimensional stability with 0% dimensional change upon hydration. The implants showed a weight increase of <20% upon hydration, indicating they are low-swelling implant devices. Furthermore, these implants could be stretched (strained) and could be stretched to >100% of their original dimensions before cohesive failure, demonstrating their elastic properties. In contrast, PLGA implants could not be stretched to a higher elongation (<10%), leading to rapid cohesive fracture. Additionally, PLGA implants became very hard upon hydration, exhibiting sharp edges.

[0437] Analysis: For applications requiring maximum contact with the underlying tissue surface, loss of dimensional stability can cause the implant to delaminate from the tissue surface. For example... Figure 33 As shown, the PCL / PTMC blend, PDO, and PCL80K maintain dimensional stability and low swelling, making these compositions suitable only for ocular use, or for any tissue requiring a flexible drug delivery system to ensure maximum comfort, such as the nose, rectum, vagina, or esophagus. PLGA is not suitable for any tissue application requiring a flexible implant / drug delivery system.

[0438] Experiment #2: In another study (Table 49 and Figure 34 In this study, a PCL blend (polycaprolactone (PC12) blended with poly(trimethylene carbonate):polycaprolactone (90:10) in a 50:50 ratio) was fabricated into a mesh insert with a thickness of ~30-50 micrometers. The insert was then incubated in phosphate buffer (pH 7.4) at 37ºC for 24 hours, followed by incubation for 72 and 14 days, respectively. The % dimensional change rate at hydration was 0%; additionally, the % hydration rate was <20%.

[0439] Table 49. Time progression study of PCL blend inserts

[0440]

[0441] In another study, PDLLA-TMC (50:50) (Sigma Aldrich) was fabricated into a coaxial mesh implant with a thickness of ~30-50 micrometers and incubated in phosphate buffer (pH 7.4) at 37ºC for 9 days. Measurements were taken at T=0 and T=9 days. Figure 35 This indicates that dimensional stability of the implant is an important product characteristic. After 9 days, the PDLLA-TMC implant lost its flexibility and shrank several times (>20% of its original size), becoming a hard implant with sharp edges, unsuitable for use in soft tissue spaces. The coaxial mesh implant contains protein (IgG) and protective excipients (phosphate pH 7, dextran, sorbitol, PEG40 stearate, cetyl alcohol) in a hydrophilic inner layer, and contains PDLLA-TMC as a hydrophobic encapsulation layer.

[0442] Table 50. Dimensional Study of PDLLA-TMC Coaxial Insertion Components

[0443]

[0444] Suitable polymers for maintaining dimensional stability are poly(trimethylene carbonate):poly(caprolactone) copolymers with a molecular weight of 80K-150K, in copolymerization ratios of 90:10, 80:20, or 70:30. Polydioxanone / polytrimethylene carbonate / polycarbonate blends are also suitable, using various ratios to adjust drug release. Another polymer that can be added is vinyl polyvinyl acetate (EVA), either added to blends with polycaprolactone and poly(trimethylene carbonate) or poly(dioxanone), or used alone.

[0445] Experiment #3: EVA Size Study

[0446] Table 51. Dimensional Study of Ethylene Acetate-Based Inserts

[0447]

[0448] The ethylene vinyl acetate inserts exhibit high dimensional stability (% dimensional change <10%, % hydration <100%). Furthermore, the ethylene vinyl acetate inserts show an elongation strain percentage >100%, indicating high elasticity. (% elongation: 20%, T=0 2.5 cm, T=3.0 cm after stretching)

[0449] Example 22: Development of Bioadhesive Implants

[0450] To impart tissue adhesion to the device, the drug-containing device may also contain water-soluble fibers with adhesive properties co-deposited onto the device, wherein the water-soluble fibers are selected from the group consisting of polyacrylic acid, polyethyleneimine, polyamide amine, polylysine, polyarginine, glucuronide chitosan, and their derivatives and blends. For use on mucosal tissues, the drug delivery system must be bioadhesive. In one example, the bioadhesive molecule is polyacrylate, an anionic bioadhesive obtained by co-spinning with a blend of poly(caprolactone) and poly(caprolactone-co-poly(trimethylene carbonate) copolymer (90:10). In other embodiments, the bioadhesive molecule may be multi-cationic, such as polyethyleneimine, polyamide amine, polylysine, polyarginine, glucuronide chitosan. A qualitative rating scale was developed to score the bioadhesive effectiveness of the implant when placed on fresh bovine conjunctiva. The rating scale was designed to assess the characteristics of adhesion and appearance. The rating scale was used to assess three different types of implants. The adhesion and appearance of polyacrylic acid (PAA) in molecular weights upon placement on conjunctival tissue were scored. Implants containing 1.8 kDa, 450 kDa, and 3 million Daltons of PAA, along with other polymer blends, were tested to determine which formulation produced ideal adhesion to the bovine conjunctiva. A PAA blend of 1.8 kDa and 450 kDa spun into a web at a ratio of 1:2 produced ideal properties for implant adhesion to the bovine conjunctival tissue. At 0, 24, and 48 hours, the implant did not slip on the tissue and exhibited resistance upon peeling, resulting in tissue stretching.

[0451] Characteristics of the test method

[0452] A qualitative rating scale was developed for the bioadhesion effectiveness of the NanoM implants when used on the conjunctival region of bovine eyelid tissue. The following criteria were used in designing the adhesion rating scale: (a) implant adhesion to the conjunctiva, (b) likelihood of implant sliding on the conjunctiva, (c) stretching of the conjunctiva after implant removal, (d) appearance of the implant on the conjunctiva, (e) likelihood of implant peeling, and (f) implant breakage during removal from the conjunctiva. Implant samples were prepared using a sterile medical-grade punch (Figure 36). Each piece was carefully “punched” from various sections of an electrospun fiber mesh. The dimensions of each implant were carefully measured and recorded in a notebook or spreadsheet. Initially, the implants were cut into square or rectangular pieces using medical-grade scissors. Fresh bovine conjunctival tissue was received on the analysis day. Upon receipt of the tissue, it was placed on ice in a refrigerator without any buffer solution until ready to begin the experiment. After the implant samples are prepared, fresh conjunctival tissue is removed from the refrigerator and placed on a sheet of wax paper. Each tissue sample can be used for more than one test, and includes both the upper and lower eyelids, each containing a segment of conjunctival tissue.

[0453] Different molecular weights of PAA spun into PCL+PCL / PTMC blends and different blend compositions resulted in varying adhesion and appearance scores. Video recording began before the implant was placed onto the tissue. In the video, the batch number, time point, and PAA content of the mesh were verbally described. Square or round implants were carefully placed onto freshly prepared conjunctival tissue using tweezers. Each tissue sample was kept on ice at room temperature until the test time, and testing was performed immediately upon receipt of the tissue to ensure freshness. During video recording, verbal descriptions of any changes in the appearance of the mesh, as well as verbal analyses of adhesion performance, were recorded. Adhesion scores were assigned at most one minute, or three days in cases involving other time points, by the presence or absence of resistance when removing the mesh from the tissue using tweezers. Adhesion scores increased when the mesh was difficult to peel from the conjunctiva. Figure 37 The ratings for adhesion and appearance can be found in Table 52 below. An adhesion score of 1 or higher at t0 meets the requirements for adhesion of the mesh at 24 hours (t24).

[0454] Table 52. Rating Levels

[0455]

[0456] Description of test samples for bioadhesion

[0457] Table 53 below summarizes the bioadhesion tests of implants made from a 50:50 blend of polycaprolactone and poly(trimethylene carbonate:polycaprolactone 90:10) co-spun with polyacrylic acid as a bioadhesive. As shown in Table 55, polyacrylic acid comprised approximately 30% of the dry implant. The polyacrylic acid ranged from 1.8 K to 450,000 K and was tested in various ratios of PAA (1.8 K) and PAA (450,000 K) at 1:1, 1:2, and 1:3. The 1:2 and 1:3 ratios yielded the highest bioadhesive fractions. Implants containing a single molecular weight of poly(acrylate) did not exhibit the strong adhesion shown in blends of PAA with varying molecular weights. This was an unexpected result, as all polyacrylates are known to have bioadhesive properties regardless of molecular weight. In contrast, implants without polyacrylic acid did not show significant tissue adhesion.

[0458] Other polymers present in dry implants that can provide bioadhesion include guar gum (1-30%), alginate (1-30%), polylysine (1-30%), polyarginine (1-30%), xanthan gum (1-30%), chitosan (1-10%), and polyethyleneimine (1-10%). In another experiment, polyarginine was incorporated into implant samples formed by electrospinning of a 50:50 blend of polycaprolactone and poly(trimethylene carbonate-co-polycaprolactone 90:10) (1-10% polyarginine). Upon hydration, the implants exhibited high adhesion (fraction = 3). Similarly, polylysine (1-10%) incorporated into the implant samples of the above blends also produced strong adhesion with a fraction of 3.

[0459] Table 53. Overview of bioadhesion studies on 50:50 blends containing PCL and PCL-P(TMC) (10:90)

[0460]

[0461] All of the implants described above are biodegradable, although their rate of degradation may be slower than the drug release rate. Therefore, these implants may need to be removed to accommodate the next implant. In another experiment, ethylene vinyl acetate (EVA) was used as the hydrophobic polymer for encapsulation, and polyarginine was used as the adhesive polymer. The adhesion score was 2–3.

[0462] Examples of "sandwich" inserts

[0463] Sandwich-type implants: These are compositions of electrospun fiber implants that focus on: (a) encapsulation of highly water-soluble and highly unstable molecules at a concentration level, (b) a method of releasing said molecules in a sustained manner, (c) minimizing irritation at the local tissue site, and (d) maximizing residence at the delivery site.

[0464] Sandwich implants have a hydrophobic drug release barrier, which takes the form of a hydrophobic "skin" surrounding a hydrophilic "core" containing a drug. To achieve a linear release rate, sandwich implants have been described in which drug-containing fibers (core-skin type or integral type) are sandwiched between two membrane layers. The described implants have a hydrophobic drug release barrier, which takes the form of a hydrophobic "skin" surrounding a hydrophilic "core" containing a drug. To achieve a linear release rate, sandwich implants have been described in which drug-containing fibers (core-skin type or integral type) are sandwiched between two membrane layers.

[0465] Without a cortex (top and bottom layers), the implant exhibited a high % protein burst release rate. This implies that core-cortex type fibers containing protein in the core do not adequately provide a barrier for sustained protein release. Although Figure 30 (Sustained release of IgG) does indeed demonstrate sustained release, but achieving this result in a reproducible manner is difficult. Typically, the % protein burst release rate is >70%, with 100% released within 1 day. To address this issue, a sandwich-type insert was designed and implemented, in which the core-shell type fibers (containing protein) are further encapsulated by two outer layers. The top and bottom layers need to be composed of "flowing" polymers; poly(trimethylene carbonate) based polymers are known to have flowability to completely encapsulate the inner core-shell type fibers. To improve the elastic properties of the poly(trimethylene carbonate) based fibers (both the core-shell type and the encapsulation), ethylene vinyl acetate can be added to the formulation. In one example, the addition of ethylene vinyl acetate to the formulation increased the elastic elongation to greater than 500%.

[0466] The selection of polymers for the coating layers (top and bottom layers) and the post-processing conditions are crucial for achieving the design characteristics of sandwich implants. Only polymers with molecular fluidity at temperatures between 28–36°C can be used to produce the coating layers. The advantage of using electrospinning instead of solvent casting is that the latter results in high concentrations of organic solvents in the drug delivery device, which degrade proteins. The electrospinning process evaporates volatile organic solvents before device fabrication. After the electrospun implant is fabricated, the polymers constituting the layers flow to create a continuous top and bottom layer.

[0467] Example 23: Implant batch number: AB-NANOM-X-07-65

[0468] Core: The core is an aqueous solution containing buffers, sugars, surfactants (Tween 20, PEG, poloxamer), salts, and proteins. All excipients help reduce surface tension and protect the proteins from degradation. The preparation of the core solution is explained step-by-step below:

[0469] Preparation of histidine buffer: To prepare histidine buffer, use a sterile container and weigh the required amount of WFI. Weigh L-histidine and L-histidine HCl using an analytical balance and add them to the container. Add a magnetic stir bar to the container and place it on a magnetic stirring plate. After the excipients are in the solution, measure the pH. Target pH: 6.0–6.4. Adjust as needed. See Table 54.

[0470] Preparation of the excipient stock solution: To prepare the excipient stock solution, weigh the required amount of WFI and then histidine buffer into a new sterile container. Add a magnetic stir bar to the container and place it on a magnetic stirring plate. Weigh all excipients except HpβCD / cyclodextrin using an analytical balance. Add all weighed excipients to the sterile container. After all components have been properly mixed, weigh HpβCD / cyclodextrin and add it to the container. Mix the solution appropriately until it becomes clear. Filter the excipient stock solution into a new sterile container using a sterile 0.22 µm PVDF syringe filter. See Table 55.

[0471] Preparation of the final core solution: To prepare the core solution, dissolve the protein (lyophilized protein stored in glass vials) directly using the excipient stock solution. Add the target amount of excipient stock solution to the glass vial and agitate to ensure complete dissolution (do not vortex). After the protein is dissolved, remove the solution from the vial and transfer it to different vials. Repeat this step for all glass vials containing protein (the vials are numbered according to the target concentration). After all proteins are fully dissolved, transfer the solution to 5 mL sterile centrifuge tubes and store them at 5°C. See Table 56 for composition.

[0472] Table 54. Preparation of Histidine Buffer (SS-NANOM-X-11-64)

[0473]

[0474] Table 55. Excipient Stock Solution (SS-NANOM-X-11-64)

[0475]

[0476] Table 56. Core Solution (SS-NANOM-X-11-64)

[0477]

[0478] Core characteristics: Visual appearance: clear or slightly turbid solution (depending on protein concentration); pH: target 6.0-6.4; Conductivity: <2.5 mS / cm; Surface tension: 30-34 mN / m

[0479] Core Characterization: Surface tension is an important parameter for the core solution as it determines the success of coaxial electrospinning. Surface tension is measured using a tensiometer (Biolin Scientific Attention Theta Lite). pH is crucial for determining protein stability; outside the range, it can lead to protein degradation. A pH meter (Fisherbrand Accumet AB315 or equivalent) is used for accurate measurement. Depending on the sample volume, pH test strips may also be used. Conductivity is another parameter determining the success of the electrospinning process. A Mettler Toledo AG Five Go or equivalent can be used; the core solution is quantified using UPLC. This determines the loading of the implant. Sheath Solution: For sandwich implants, two types of sheath solutions are used: coaxial sheaths and membrane sheaths. A coaxial sheath is a polymer solution that encapsulates the core solution (also known as the intermediate layer). Membrane sheaths are used in the top and bottom layers. These two solutions are used together to form a sandwich mesh / implant.

[0480] Preparation method: Weigh all solids for both solutions in several separate weighing boats; after weighing all solids, tare the glass vial for the first solution on an analytical balance. Draw 3 mL of DMSO from the DMSO dispensing vial using a 3 mL syringe and a 25 G needle. Weigh the required amount of DMSO into the tare vial; tare the vial again and add all the solids required for this solution. Tare the vial again. Draw at least 10 mL of methyl acetate from the sealed vial using a 10 mL syringe with a 25 G needle. Add the methyl acetate to the tare vial and record the weight. Repeat the above steps for the second solution (membrane sheath). Label the vials appropriately, listing all solvents and polymers; place both vials on a rotating rack in an incubator set to 40–42°C. Mix the solutions overnight.

[0481] Table 57. Composition of the intermediate layer of the coaxial sheath

[0482]

[0483] Table 58. Composition of the outer coating layer

[0484]

[0485] Sheath properties

[0486] Visually, the sheath is a clear solution without undissolved polymer. Viscosity was measured using an AntonPaar MCR92 rheometer (cone-plate type). Successful electrospinning requires a viscosity of 1500-4000 cP.

[0487] Table 59. Specific Concentration Logarithmic Viscosity of Primary Coated Polymers

[0488]

[0489] Table 60. Electrospinning Parameters

[0490]

[0491] For P2, the coating layer and core-sheath solution are co-spun.

[0492] Encapsulation Procedure. This procedure is used to analyze proteins / drugs encapsulated within implants. When the top and bottom layers are composed of polymers, the polymers are dissolved using an organic solvent. For this method, multiple washes are performed using a 50:50 methyl acetate:dichloromethane mixture to dissolve the polymers. After the washing process, the protein pellets are dried to remove residual solvent. Following this step, the dried pellets are reconstituted using 50 mM phosphate buffer, 1.2% sodium chloride, and 0.1% poloxamer. The implants are analyzed using UPLC. Sampling of the encapsulation mesh: Using a sharp, sterile 10 mm skin piercing tool, remove one punch from the center of the mesh, avoiding the stable area; place the sample in a peeled 1.5 mL centrifuge tube and record the sample weight (the sample weight should be ~10 mg). Separate the polymeric components from the aqueous components: Using a syringe and needle, draw 0.4 mL of dichloromethane and 0.4 mL of methyl acetate from their respective vials and add them to centrifuge tubes; vortex the centrifuge tubes 2-3 times at 5-second intervals to break the implant into small pieces; after the implant is broken, sonicate the tubes at 5-second intervals for 1-1.5 min or until the implant is almost completely dispersed; after the implant is completely dissolved / dispersed, place the tubes in a temperature-controlled benchtop centrifuge (@ 13.2K rpm and 5°C) for 30 minutes.

[0493] Protein Washing Procedure: Remove the sample from the centrifuge and observe its appearance. The suspended material will form pellets, and the solvent phase should be clear. Use a 1000 µL pipette to remove ~0.6–0.8 mL of solvent, being careful not to disturb the pellets or any small insoluble particles. Remove the solvent to remove the sheath component; repeat steps c through h three more times, adding fresh solvent at each step. Skip step e, as sonication is only required for the initial dissolution. Drying the Protein Pellets: Open the centrifuge tube cap and seal the tube tightly with sealing film. Use a sterile needle (26–30 G) to prick 3–4 holes in the sealing film. The holes should be large enough to allow solvent vapor to escape during the drying process. Dry the pellets using a CentriVap connected to a cold trap. Place the centrifuge tube inside the CentriVap and evaporate the residual solvent at 25°C for 30 minutes. After the drying cycle is complete, visually inspect the pellets. They should be dry, white, and intact.

[0494] Reconstitution of Protein Pellets: Note that the protein load in the implant and the weight of the original dry implant sample are used together to determine how much buffer should be used to reconstitute the pellets. The pellets should be reconstituted to the concentration within the standard curve. Carefully remove the sealing film from the centrifuge tube and reconstitute the pellets using a calculated amount of 50 mM phosphate buffer and 1.2% sodium chloride. Mix the solution by vortexing 2–3 times at 5-second intervals. Incubate the sample at room temperature for 40 min to ensure complete decomplexation of the protein with the residual sheath polymer. A very small amount of white precipitate may be present in the solution after pellet dissolution, for example, in the presence of residual sheath polymer. Centrifuge the solution for 15 sec using the short spin option to separate the buffer solution from the precipitate. Take 100 µL of the clear phase using a 200 µL pipette and transfer it to a small-volume polypropylene HPLC vial. Determine the sample and freshly prepared STD curve using HPLC for protein analysis. Examples 21 and 22 are examples of sandwich inserts having proteins with MW of <20K and 60-70K.

[0495] When sandwich inserts are not used, the incorporated proteins are released almost immediately. Figure 39 This demonstrates the necessity of sandwich-type inserts for achieving sustained protein release, with a burst release rate of <30%. Figure 38 The appearance of a core-skin type coaxial insert without a sandwich microstructure is shown. The composition of the core-skin type (intermediate) layer is the same as that of the sandwich type insert. Figure 38 The insert in the middle shows electrospun fibers, rather than a sandwich insert. Figure 41 The unique microstructure shown in )

[0496] Figure 39 The implant, which only contains a core-skin type intermediate layer and no overlay, did not continuously release.

[0497] For protein-incorporated core-sheath fibers, the coaxial microstructure of the fibers is insufficient to provide for the release of bioactive compounds. When encapsulated in a "sandwich" form, sustained release can be achieved.

[0498] Example 24: Sandwich (Layered) Insert Incorporating Protein (MW < 25K) and Diffusion Aid

[0499]

[0500]

[0501] Example 25: Sandwich (Layered) Insert Incorporating Protein (MW < 70K) and Diffusion Aid

[0502] Table 63. AB-NANOM-03-56 Insert Components

[0503]

[0504]

[0505] Example 26: Sandwich (Layered) Insertion Incorporating Protein (MW>100K) and Diffusion Aid PTMC(3k)-PEG(3.4k)-PTMC(3k)

[0506]

[0507] Example 27: Sandwich (layered) insert incorporating protein (MW<70K) and diffusion aid (containing PTMC homopolymer with a specific logarithmic viscosity of 0.65 dL / g)

[0508]

[0509]

[0510] Example 28: Sandwich (layered) implant incorporating protein (MW<70K) and diffusion aid (containing PCL-PTMC (10:90) and PCL-PTMC (90:10) in the outer layer)

[0511]

[0512] External release mode of the "sandwich-type" implant

[0513] Procedure: 1. Punch out a 10 mm slurry and record its weight. 2. Take a polypropylene vial and record the tare weight of the empty vial. 3. Place the insert into the vial. 4. Add 50 µL of phosphate buffer to the vial and record its weight. 5. Place the vial in an incubator set to 37°C. 6. After 30 min, remove the vial from the incubator. 7. Record the total weight of the vial. 8. Take 20–30 µL of buffer and add it to the HPLC vial. 9. Record the weight of the taken buffer. 10–30 µL of the above buffer is replenished with freshly prepared phosphate buffer containing sorbitol, and the vial is returned to the incubator. 10. Repeat the above steps for each time point (4 hours, 1 day, 2 days, etc.). 11. Dilute the buffer recovered at each time point with 20 µL of phosphate buffer. 12. Perform HPLC on the sample and analyze the released proteins using UPLC.

[0514] Figure 40This demonstrates a sustained release mode for the "sandwich" implant. Sustained release of highly water-soluble proteins is not feasible without the unique microstructure of the sandwich implant. The flowable polymer consists of poly(trimethylene carbonate), copolymerized with other polymers such as poly(caprolactone), or blended with other polymers used to impart strength. The molecular weight of the poly(trimethylene carbonate) will affect the flowability of the encapsulated implant. For example, if the specific logarithmic viscosity of poly(TMC)-co-PCL is 1.2 dl / g, a partial encapsulation layer is formed after preparation without prolonged annealing (30-35°C). Furthermore, continuity is further established when the implant is sterilized by irradiation at 6 kGy.

[0515] The "intermediate layer" remains in the form of electrospun fibers. Figure 42 ).like Figure 41 As shown, the coating layer initially appears as an electrospun fiber, but transforms into a continuous layer (top and bottom layers) as the coating. The coating polymer is essential to slow down the incorporation of bioactive agents. The diffusion aid poly(trimethylene carbonate)-polyethylene oxide-poly(trimethylene carbonate) serves to generate a water-soluble phase that allows the incorporated protein to diffuse through.

[0516] The polymers in the coating layer need to be able to “flow” to transform the top layer into a continuous layer. These polymers are based on poly(trimethylene carbonate) polymers as well as blends and copolymers with poly(caprolactone) for increased strength and durability.

[0517] In one instance, when poly(trimethylene carbonate)-co-caprolactone (90:10) was used as the sole coating polymer, the resulting product flowed excessively, exposing the incorporated bioactive proteins and resulting in a protein burst release rate of ~89%. Therefore, polycaprolactone is needed to add strength and structure to the top and bottom layers of the implant. With full coating, protein burst release decreased to <25%, but further protein release could be achieved by increasing the concentration of poly(TMC)-PEG-poly(TMC) in the coating layer.

[0518] The effects of electron beam and temperature on polymer flowability.

[0519] The membrane sheath of the mesh was composed of 18% w / v (90:10) PCL-PTMC:PCL with 0.116% PTMC-PEG-PTMC (diffusion aid). This composition exhibited a favorable release pattern (20.1% burst release at 30 min and 23.4% burst release at 30 min after electron beam (6 kGy)). Figure 43SEM images before and after annealing are shown. The image at T = 24 hr shows a sharp reduction in the number of pores, indicating the effect of temperature (annealed at 33°C) on the polymer flowability. The third image shows the effect of the electron beam on the structure of the mesh. The number of pores is visually reduced slightly, but has a significant effect on the release mode. This is a good example showing the effect of temperature and electron beam on the release mode and polymer flowability.

[0520] The energy associated with electron beam radiation allows the coated polymer to flow, having a similar effect to annealing. Figure 44 The release rates of the model protein before and after annealing are shown, with the % protein burst rate decreasing from 60% (before annealing) to 35% (after annealing). Scanning electron micrographs of the annealed slices show a more continuous outer layer. On the other hand, Figure 45 The release of proteins from the implant is shown before and after electron beam sterilization. The release pattern before sterilization is a rapid release at a rate greater than 60% within 1 day, while post-sterilization exhibits a more sustained protein release. Figure 43 As shown, both annealing and the energy from the sterilization radiation have the effect of causing the outer layer to flow, thus forming a diffusion barrier against protein release. Other sterilization radiation, such as X-rays or gamma radiation, also have a similar effect to electron beams (β radiation). The sterilization radiation is in the range of 6-25 kGy, preferably at low temperatures.

[0521] Other implementation plans

[0522] All features disclosed in this specification can be combined in any way. Each feature disclosed in this specification can be replaced by alternative features for the same, equivalent, or similar purposes. Therefore, unless otherwise expressly stated, each disclosed feature is merely an instance of a general set of equivalent or similar features.

[0523] Based on the above description, those skilled in the art can readily identify the essential features of the present invention, and various changes and modifications can be made to adapt it to a variety of uses and conditions without departing from its spirit and scope. Therefore, other embodiments are also within the scope of the appended claims.

Claims

1. A dry drug delivery system comprising an intermediate layer enclosed within two outer layers, the intermediate layer and the two outer layers each comprising a mixture of electrospun fibers, the dry drug delivery system having a density of 0.1-1 g / cm³. 3 The final total density within the range, thickness <1 mm, elastic modulus <4 MPa, and elastic strain change >100%, where Each of the two outer layers comprises one or more polymers selected from the following: ethylene vinyl acetate, polytrimethylene carbonate (P(TMC), P(TMC)-co-poly(caprolactone) (PCL) in a copolymer ratio of 90:10, PCL, and P(TMC)-polyethylene oxide (PEG)-P(TMC), wherein the PEG segment has a molecular weight of 1,000-20,000 Da and the P(TMC) segment has a molecular weight of 1,000-10,000 Da; and The intermediate layer is formed of core-sheath type fibers, each of which comprises: A core, the core containing a bioactive substance and one or more excipients that protect the bioactive substance from degradation, and The outer skin is formed of a hydrophobic elastic polymer, which is composed of P(TMC) and PCL in a weight ratio of 90:10 to 30:

70.

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

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

4. The dry drug delivery system according to any one of claims 1 to 3, wherein, based on the weight of the two outer layers, the 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).

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

6. The dry drug delivery system according to any one of claims 1 to 5, wherein the bioactive substance is a water-soluble drug selected from the following: small molecules with a molecular weight <1000 Da, large molecules with a molecular weight >1000 Da, proteins, amino acids, peptides or nucleic acids with a molecular weight of 10,000 Da to 250,000 Da.

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

8. The dry drug delivery system of claim 1, wherein the bioactive substance is water-soluble and selected from the group consisting of: immunoglobulins, antigen-binding domains, antibacterial agents, antiviral agents, steroids, antiglaucoma agents, NSAIDs, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, antifungal agents, anti-inflammatory drugs, proteins, peptides, growth factors, enzymes, vitamins, hormones, polysaccharides, peptide-nucleic acids, and nucleic acids or mixtures thereof.

9. The dry drug delivery system of claim 1, wherein the bioactive substance is hydrophobic and selected from the group consisting of: antimicrobial agents, steroids, NSAIDs, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, antifungal agents, anti-inflammatory drugs, and antiglaucoma agents.

10. The dry drug delivery system according to any one of claims 1 to 9, further comprising a water-soluble adhesive fiber deposited on the system, the water-soluble adhesive fiber being selected from polyacrylic acid, PEG-polyarm amine, polyethyleneimine, polyamide amine, polylysine, polyarginine, glucuronide chitosan, and their derivatives and blends.

11. The dry drug delivery system according to any one of claims 1 to 10, wherein during the preparation of the system by electrospinning, the one or more excipients in the core of the core-shell fiber interact with the bioactive substance as a water substitute, cooperate with the bioactive substance, or form an interface layer between the polymer-containing organic phase and the aqueous phase.

12. The dry drug delivery system of claim 11, wherein the one or more excipients are selected from the group consisting of: hydroxypropyl β-cyclodextrin, γ-cyclodextrin, sulfobutyl β-cyclodextrin, PEG40 stearate, trehalose, sorbitol, sucrose, mannitol, poloxamer 407, poloxamer 188, polysorbate 80, polysorbate 20, dextran, polyvinyl alcohol, polyethylene oxide, polyvinylpyrrolidone, hydroxypropyl methylcellulose, human serum albumin, and soluble collagen, wherein the one or more excipients are present in the dry drug delivery system at a concentration of 1-75% w / w.

13. The dry drug delivery system of claim 1, wherein the system has been sterilized by low-temperature electron beam radiation with a dose range of 6-25 kGy.

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

15. The dry drug delivery system of claim 1, wherein the hydrophobic elastic polymer of the outer sheath forming the core-shell type fiber 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 the soft tissue of a subject, the method comprising: Obtain the dry drug delivery system of claim 1, and The system is applied to the surface of the soft tissue before or after adding a hydration fluid to the subject's soft tissue. The hydration fluid comprises an isotropic solution containing one or more of the following: sodium chloride, phosphate, citrate, albumin, magnesium salt, calcium salt, borate, boric acid, balanced salt solution, multi-branched PEG-amine, PVA, multi-branched polylysine, multi-branched polyarginine, guar gum, gellan gum, sodium alginate, xanthan gum, carboxymethyl cellulose, and the hydration fluid has a pH of 6-7.8 and an osmotic pressure of 270-340 mOsm / kg.

17. A water-soluble drying device, the device comprising a physical mixture of a first fiber and a second fiber different from the first fiber, The first fiber contains sodium hyaluronate, polyethylene glycol (PEG), and polysorbate 80 or polysorbate 20, and optionally contains castor oil, buffer salts, amino acids, calcium chloride, mannitol, sucrose, trehalose, sodium alginate, polyvinylpyrrolidone, dextran 70, albumin, PEG-polyamine, type I collagen, glycerol, or PEG40 stearate. The second fiber contains polyvinyl alcohol (PVA) and sodium alginate, and optionally contains sodium borate, dextran 70, PEG, PEG-multi-arm succinimide, polypropylene glycol, propylene glycol, type I collagen, bioactive substances, sodium hyaluronate, xanthan gum, guar gum, cyclodextrin, trehalose, mannitol, sorbitol, tamarind seed polysaccharide, or β1>6 glucan, wherein the first fiber and the second fiber are configured to react with each other and with biological tissue upon hydration, and the device transforms into a hydrogel upon contact with the biological tissue and subsequent hydration, and completely dissolves within 0 h to 14 days after contact with the biological tissue.

18. The water-soluble drying apparatus of claim 17, wherein the PVA has a molecular weight of 88,000-130,000 Dal and a degree of hydrolysis of 88-100%, and the sodium hyaluronate has a molecular weight of 50,000-5,000,000 Dal, and the sodium alginate has a molecular weight of 32,000-400,000 Dal, the first fiber comprising 80-90 wt% PEG, 0.5-5 wt% sodium hyaluronate, 0.1-0.3 wt% sodium chloride, 0.001-2 wt% polysorbate 20 or polysorbate 80, and the second fiber comprising 80-95 wt% PVA.

19. The water-soluble drying device according to claim 17 or 18, wherein the water-soluble drying device further comprises a bioactive substance selected from the following: small molecules with a molecular weight <1000 Da, large molecules with a molecular weight >1000 Da, proteins, amino acids, peptides or nucleic acids with a molecular weight of 10,000 Da to 250,000 Da.

20. The water-soluble drying device according to claim 19, wherein the bioactive substance is water-soluble, slightly water-soluble, or water-insoluble.

21. The water-soluble drying device according to claim 17, wherein, When placed on biological tissue, the device dissolves either through simple dissolution in an aqueous fluid or through biodegradation.

22. The water-soluble drying device according to any one of claims 17 to 21, wherein the device is circular, oval, crescent-shaped, elliptical, annular, square, rectangular, or tubular.

23. The water-soluble drying device according to claim 17, wherein the device has been sterilized by low-temperature electron beam radiation with a dose range of 6-25 kGy.

24. The water-soluble drying device according to claim 19, wherein the first fiber or the second fiber comprises the bioactive substance, the bioactive substance being selected from immunoglobulins, antigen-binding domains, antibacterial agents, antiviral agents, steroids, antiglaucoma agents, NSAIDs, ultraviolet blockers, healing agents, decongestants, antihistamines, antioxidants, vasoconstrictors, antiallergic drugs, lubricants, analgesics, antifungals, anti-inflammatory drugs, proteins, peptides, growth factors, enzymes, vitamins, hormones, polysaccharides, peptide-nucleic acid, nucleic acids, or mixtures thereof.

25. The water-soluble drying device according to claim 21, wherein the biological tissue is an ocular mucosa, vaginal mucosa, rectal mucosa, nasal mucosa, or oral mucosa.

26. A method for delivering a bioactive substance to the soft tissue of a subject, the method comprising: Obtain the water-soluble drying device as described in claim 17, and The system is applied to the surface of the subject's soft tissue. The device transforms into a hydrogel upon contact with the surface of the soft tissue through hydration.

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