Inflammation-responsive Anti-inflammatory hydrogels

Abstract

To provide protease-responsive drug delivery hydrogels, uses thereof, and related methods of their production, more particularly, to provide hydrogels which release anti-inflammatory agents upon reaction with inflammation-related proteases.SOLUTION: A drug-loaded protease-responsive hydrogel is provided, comprising: a) a drug covalently bound to a protease-cleavable peptide anchor having a functional moiety; b) a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional moiety; and c) a bis-functional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional moieties; wherein the drug and the arm of the multi-arm PEG polymer are covalently bound through the functional moiety of the peptide anchor, and a gel is formed by covalent bonding between the functional moiety of the polymer building block and the functional moiety of the bis-functional crosslinker.SELECTED DRAWING: Figure 19

Description

[Technical Field]

[0001] The present invention relates generally to the field of protease-responsive drug delivery hydrogels, their uses, and methods related to their manufacture, and more particularly to hydrogels that release anti-inflammatory drugs upon reaction with inflammation-associated proteases. [Background technology]

[0002] Inflammation is a series of biological responses initiated by the host immune system to remove harmful stimuli and restore damaged tissues to their pre-injury state [Serhan, CN et al. Fundamentals of Inflammation, Cambridge University Press, Cambridge, (2010)]. Acute inflammatory responses are essential for eliminating harmful stimuli and restoring cellular homeostasis after tissue injury [Serhan, CN et al. Fundamentals of Inflammation, Cambridge University Press, Cambridge, (2010)]. However, excessive leukocyte activation can lead to chronic tissue damage due to excessive inflammation [Serhan, CN et al. Fundamentals of Inflammation, Cambridge University Press, Cambridge, (2010)]. Such chronic conditions are common in many conditions, including rheumatoid arthritis, chronic diabetic ulcers, inflammatory bowel disease (IBD), and chronic obstructive pulmonary disease [Serhan, CN et al. Fundamentals of Inflammation, Cambridge University Press, Cambridge, (2010)].

[0003] Systemic administration of anti-inflammatory drugs has been clinically approved as a treatment for suppressing excessive inflammation in chronic diseases. Patients with rheumatoid arthritis and IBD are empirically prescribed small-molecule drugs, such as nonsteroidal anti-inflammatory drugs (NSAIDs) and steroid immunosuppressants, based on clinical symptoms. However, systemic administration of these drugs can lead to overdosing due to uncontrolled drug release, which can result in well-known side effects. For example, systemic administration of NSAIDs increases the risk of myocardial infarction, cerebrovascular accident, and gastric ulcers. Furthermore, long-term prescription of corticosteroids can lead to serious drug-induced complications, such as osteonecrosis, glaucoma, and opportunistic infections.

[0004] Several drug delivery platforms for anti-inflammatory therapeutics have been designed to improve spatiotemporal control of drug release kinetics and minimize systemic toxicity [Hamalainen, M. et al. Basic & Clinical Pharmacology & Toxicology 112(5) 296-301 (2013)]. For example, encapsulation of glucocorticoids in a vesicular system has been reported to extend the drug's half-life and achieve slow release [Maestrelli, F. et al. Journal of Drug Delivery Science and Technology 32 192-205 (2016)]. Additionally, covalent attachment of small NSAIDs to nanosized polymer films significantly increased the therapeutic payload, resulting in sustained release over a long period of time via hydrolysis of the drug-polymer ester bond [Hsu, B.B. Proceedings of the National Academy of Sciences 111(33) 12175 (2014)]. However, these systems do not take into account the specific pathology of diseased tissues with inflammatory characteristics that require the administration of anti-inflammatory therapeutic agents.Drug release is primarily determined by the physicochemical properties of the delivery platform, such as polymer composition and drug loading, and therefore the drug release profiles of these systems do not match biological requirements.

[0005] By exploiting the inflammatory characteristics of the biological microenvironment in diseased tissues, it may be possible to design smart drug delivery systems that are driven by immunological signals. In particular, several studies have shown that chronic inflammation increases the expression of proteases, particularly serine proteases and matrix metalloproteinases (MMPs), suggesting that these may serve as biochemical signals for therapeutic administration to modulate the inflammatory cascade [Pham, CTN The International Journal of Biochemistry & Cell Biology 40(6) 1317-1333 (2008)]. Proteases are more specific biological signals than other stimuli, such as pH, temperature, and redox, primarily because dysregulation of proteases is closely linked to pathology. Furthermore, other stimuli may be affected by environmental conditions. For example, body temperature can rise rapidly in humid weather conditions, even in the absence of disease. Despite the advantages of proteases as biological signals, their potential as immunological signals for biologically driven drug delivery systems to modulate inflammation has been largely unexplored. Recently, Joshi et al. utilized the self-assembly properties of the amphiphilic small molecule triglycerol monostearate (TG-18) to physically encapsulate a corticosteroid within a hydrogel platform and successfully induce the release of the drug in response to reactivation of arthritis activity [Joshi, N. et al. Nature Communications 9(1) 1275 (2018)]. However, the reported drug-loaded hydrogel lacks a generalizable design framework, limiting the components that can be modified to exploit alternative biological factors.Specifically, drug release from this platform relies primarily on cleavage of the ester bond in the TG-18 backbone by esterases, specifically, several esterases whose expression is elevated in inflammatory arthritis [Ravaud, P. et al. Rheumatology 41(7) 815-818 (2002)]. However, these esterases may not be an important biological signal in other inflammatory diseases. Furthermore, nonenzymatic hydrolysis of this ester bond may occur in the low pH environment associated with inflammatory conditions [Bellocq, A., et al. Journal of Biological Chemistry 273(9) 5086-5092 (1998); Riemann, A. et al. Molecular Basis of Disease 1862(1) 72-81 (2016)], potentially resulting in undesired nonspecific drug release.

[0006] Therefore, there is an unmet need for a protease-triggerable drug delivery platform that addresses the limitations of existing delivery systems: (1) modular in design, (2) immunocompatible, and (3) versatile enough for both injectable and topical administration at room temperature. First, drugs can be physically encapsulated within particulate domains, such as liposomal or polymeric microparticles, which can then be incorporated into protease-triggerable delivery systems, resulting in a basal release of the drug via diffusion. While such a basal release may be desirable for the management of chronic inflammatory diseases that require protease-triggered dosing during acute exacerbations such as infections or arthritis flares, it is not desirable for all inflammation-related conditions, particularly in immunocompromised or immunosuppressive patients, or for the management of acute injuries where a certain level of inflammation is required for normal healing. For conditions in which the drug administration site transitions from a drug-free physiological state to a highly inflammatory pathological state, such as the sudden onset of bacterial infection at an acute wound site or a flare-up of seborrheic dermatitis, alternative designs of protease-inducible delivery systems that eliminate or minimize basal drug release are desirable.

[0007] Furthermore, in the management of inflammation-related pathologies, using a single protease as a biochemical stimulus to induce drug release can sometimes allow for the dosage to be tailored to the inflammatory state of the disease. However, the expression of multiple proteases may be elevated in pathological inflammatory conditions. Therefore, by using multiple proteases rather than a single protease, it is possible to enhance the specific correlation between protease activity and each disease state, which may enable drug release kinetics tailored specifically to the target inflammation-related disease. Therefore, there remains a great need for the development of drug delivery systems that release drugs in response to stimulation by two or more proteases (or in response to multiple proteases) with the aim of improving specificity. Summary of the Invention [Means for solving the problem]

[0008] The present invention provides an inflammation-responsive drug delivery platform comprising: (1) a drug-loaded domain (particles or conjugates with anti-inflammatory drugs) with a tailored basal release profile of the drug, and / or (2) Protease-cleavable hydrogel domain The present invention provides an inflammation-responsive drug delivery platform comprising: The present invention provides a drug delivery platform that can be customized to address inflammatory diseases, and its responsiveness and specificity can be adjusted to suit the target disease by changing the composition of the drug-carrying domain and / or adjusting the sensitivity of the protease-responsive domain to multiple types of proteases.

[0009] According to a first aspect of the present invention, there is provided a protease-responsive drug-loaded hydrogel, comprising: a) Drugs encapsulated in particles; b) a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences containing functional groups; Including, A protease-responsive drug-loaded hydrogel is provided, characterized in that the polymer building block (b) forms a gel in the presence of the protease-cleavable crosslinker (c), and the particles (a) are encapsulated in the gel.

[0010] In some embodiments, the protease-responsive drug-loaded hydrogel comprises: a) at least one second bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by spacer sequences comprising functional groups; and / or b) at least one bifunctional protease-resistant crosslinker containing a protease-resistant substrate; Further including, The protease-cleavable substrate of a) is characterized by being sensitive to a protease different from the protease to which the cross-linking agent of c) is sensitive.

[0011] The particles may be composed of any material capable of loading and releasing drugs (such as small molecules, therapeutic peptides, proteins, or mRNA) and encapsulated in a gel formed by the polymer building blocks and the crosslinker. The particles may be, for example, silica, liposomes, siRNA complexes, or polymeric materials. The particles can be manufactured using well-known prior art methods, such as emulsion, electrospray, electrostatic complexation, and flow focusing [Abdelaziza, Hadeer M. et al., Journal of Controlled Release 269 374-392 (2018)].

[0012] In some embodiments, the drug is encapsulated in particles comprising a polymeric material selected from the group consisting of polycaprolactone, poly(methacrylic acid), polylactic acid, polyvinylpyrrolidone, poly(lactic-co-glycolic acid) (PLGA), and gelatin. The particles are preferably microparticles and / or nanoparticles, and the particle diameters preferably range from about 10 nm to about 100 μm.

[0013] In some embodiments, the polymer building block comprises a multi-arm PEG-vinyl sulfone, a multi-arm PEG-maleimide, a multi-arm PEG-azide, or a multi-arm PEG-alkyne. Sulfone groups are advantageous because they interact with cysteine ​​groups on the arms of the crosslinker.

[0014] The present invention also embodies protease-responsive drug-loaded hydrogels in which the drug does not need to be encapsulated in particles and kept encapsulated until released by the protease.

[0015] According to a second aspect of the present invention, there is provided a protease-responsive drug-loaded hydrogel, comprising: a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, A protease-responsive drug-loaded hydrogel is provided, characterized in that the drug and the arms of the multi-arm PEG polymer are covalently bonded via the functional groups of the peptide anchor, and a gel is formed by covalently bonding the functional groups of the polymer building block and the functional groups of the bifunctional crosslinker.

[0016] Depending on the arrangement of peptide anchors and crosslinkers, the release profile of drug-loaded hydrogels can be flexible and tunable, allowing for the release of drugs sensitive to one or more different proteases.

[0017] Advantageously, the drug binding domain of this embodiment minimizes the basal release of the drug.

[0018] In some embodiments, the protease-responsive drug-loaded hydrogel comprises: a) the cross-linker is not protease-cleavable; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease. It is characterized by:

[0019] The desired peptide anchor consists of a protease-cleavable spacer sequence containing a functional group, advantageously the spacer sequence containing at least four amino acids.

[0020] Advantageously, the cross-linker comprises a sequence of a protease-cleavable substrate flanked by spacer sequences containing functional groups, each spacer sequence containing at least four amino acids.

[0021] The non-cleavable crosslinker is used to control the diffusion of the protease into the gel network, which is useful for tailoring the release profile.

[0022] In some embodiments, the drug may be a small molecule, an siRNA, an aptamer, or a therapeutic peptide or protein.

[0023] Combining multiple peptide sequences, which are the main components of the protease-responsive domain, is advantageous because it rapidly forms an aqueous gel and improves the specificity of induced release upon contact with multiple disease-specific proteases.

[0024] In some embodiments, the polymer building block comprises a multi-arm PEG-vinylmaleimide. The amount of drug loaded into the protease-responsive hydrogel can be adjusted by the amount or concentration of the multi-arm PEG polymer used.

[0025] In some embodiments, the weight ratio of the protease-responsive drug-loaded hydrogel is about 2 w / v% to about 12 w / v%, preferably about 3 w / v% to about 10 w / v%, and the hydrogel is preferably multi-arm PEG-vinyl sulfone, multi-arm PEG-vinyl maleimide, multi-arm PEG-alkyne, or multi-arm PEG-azide.

[0026] It will be appreciated that the number of arms in the multi-arm PEG polymer will affect the amount of drug that can be bound, as well as the degree of cross-linking and gel formation.

[0027] In some embodiments of the protease-responsive drug-loaded hydrogel according to any aspect of the present invention, the multi-arm PEG polymer has 3 to 8 arms.

[0028] In some embodiments of the protease-responsive drug-loaded hydrogel according to any aspect of the present invention, the drug is an anti-inflammatory agent.

[0029] In some embodiments of the protease-responsive drug-loaded hydrogel according to any aspect of the present invention, the protease is a protease whose expression increases during inflammation and is selected from the group including matrix metalloproteases and serine proteases.

[0030] In some embodiments of the protease-responsive drug-loaded hydrogel of any aspect of the present invention, the drug is a steroidal anti-inflammatory drug, a nonsteroidal anti-inflammatory drug (NSAID), or a derivative thereof. The drug may be a steroidal anti-inflammatory drug such as dexamethasone, fludrocortisone, methylprednisolone, prednisolone, prednisone, hydrocortisone, or a derivative thereof. Glucocorticoids can be oxidized to add a carboxyl functional group, and once converted, they can be conjugated to the peptide anchor of the present invention. The drug is preferably an NSAID, such as ibuprofen, ketoprofen, dichlorofenac, sanlindac, piroxicam, celecoxib, or a derivative thereof.

[0031] In some embodiments of the protease-responsive drug-loaded hydrogel according to any aspect of the present invention, the spacer sequences flanking the substrate comprise at least one cysteine ​​residue, at least one lysine residue, and / or at least one azide- or alkyne-containing non-natural amino acid. These are necessary for the crosslinker to react with the functional groups of the multi-arm PEG to form a gel. The spacer may have 1 to 6 amino acids. The remaining residues may be any amino acid, preferably an amino acid with a charged side chain. Specifically, a positively charged amino acid (e.g., arginine, R) near the thiol group of a cysteine ​​increases the crosslinking rate, while a negatively charged amino acid (e.g., aspartic acid, D) slows the crosslinking reaction. The spacer may have 1 to 6 amino acids. In some embodiments, the spacer sequences flanking the substrate ("SPACER") may have a sequence represented by the formula GX1X2X3 (SEQ ID NO: 33) and / or its inverse. wherein X1, X2, and X3 are each independently glycine, cysteine, aspartic acid, or arginine. In some embodiments, the spacer sequences flanking the substrate are selected from the group consisting of GRCR (SEQ ID NO: 1), GCRG (SEQ ID NO: 2), GRCD (SEQ ID NO: 3), GCDR (SEQ ID NO: 4), GCDG (SEQ ID NO: 5), GDCD (SEQ ID NO: 6), GCDD (SEQ ID NO: 7), GCRD (SEQ ID NO: 8), and GCRR (SEQ ID NO: 9).

[0032] When a first and a second spacer are used, one at each end of the peptide substrate, the second spacer sequence may be the reverse of the first spacer sequence, and may be a sequence represented by the formula X3X2X1G (SEQ ID NO: 34). This reverse spacer sequence, hereinafter sometimes referred to as "RECAPS," may be, for example, the reverse sequence of a spacer selected from the group consisting of SEQ ID NO: 1, SEQ ID NO: 2, SEQ ID NO: 3, SEQ ID NO: 4, SEQ ID NO: 5, SEQ ID NO: 6, SEQ ID NO: 7, SEQ ID NO: 8, and SEQ ID NO: 9.

[0033] In some embodiments of the protease-responsive drug-loaded hydrogel according to any aspect of the present invention, the protease-cleavable substrate is sensitive to a protease selected from the group including matrix metalloproteases such as metalloprotease-9 (MMP-9), MMP-2, MMP-7, and MMP-12; cathepsins such as cathepsin K, cathepsin B, and cathepsin S; human neutrophil elastase (HNE); caspases; and urokinase.

[0034] In some embodiments, the protease-cleavable substrate is selected from the group consisting of an MMP-9 substrate comprising the amino acid sequence represented by KGPRSLSGK (SEQ ID NO: 30), GPRSLSG (SEQ ID NO: 10), LGRMGLPGK (SEQ ID NO: 11), AVRWLLTA (SEQ ID NO: 12), or GPQGIWGQ (SEQ ID NO: 13), an HNE substrate comprising APEEIMDRQ (SEQ ID NO: 14) or PMAVVQSVP (SEQ ID NO: 15), a cathepsin B substrate comprising GRRGLG (SEQ ID NO: 16) or DGFLGDD (SEQ ID NO: 17), and combinations thereof.

[0035] A third aspect of the present invention provides a composition comprising the protease-responsive drug-loaded hydrogel of any of the aspects of the present invention, formulated for injection or topical administration.

[0036] The inflammation-responsive drug-loaded hydrogel of any of the aspects of the present invention can be incorporated into a polymer dressing to form a composite dressing for wound management.

[0037] A fourth aspect of the present invention provides a dressing material comprising the protease-responsive drug-loaded hydrogel of any of the aspects of the present invention.

[0038] According to a fifth aspect of the present invention, the protease-responsive drug-loaded hydrogel according to any aspect of the present invention or the composition of the present invention is provided as an injection or an external dressing for treating a subject in need thereof.

[0039] According to a sixth aspect of the present invention, there is provided a method of treatment comprising administering an effective amount of a protease-responsive drug-loaded hydrogel of any aspect of the present invention or a composition of the present invention to a subject in need of such treatment. In some embodiments, the administration is by injection or topical application to the subject. In some embodiments, the treatment is for an inflammation-related condition in which inflammation management is desired, such as chronic wounds, inflammatory bowel disease, arthritis, or a latent infection-related condition.

[0040] According to a seventh aspect of the present invention, a) Drugs encapsulated in particles; b) a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences containing functional groups; Including, a kit, characterized in that a) to c) are as defined in any of the preceding aspects; or a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm PEG polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, a) to c) are the same as those defined in any of the above aspects. is provided.

[0041] In some embodiments, the kit comprises a protease-responsive drug-loaded hydrogel according to any aspect of the present invention or a composition according to any aspect of the present invention.

[0042] According to an eighth aspect of the present invention, there is provided a method for producing a protease-responsive drug-loaded hydrogel, comprising the steps of: a) mixing a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG) with drug-loaded particles; b) mixing the drug-loaded polymer particles with a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by spacer sequences comprising functional groups; and c) mixing the mixture of a) with the mixture of b) so that the polymer building blocks of a) form a gel in the presence of the protease-cleavable crosslinker of b), and the drug-loaded particles are encapsulated in the gel. A method of manufacturing is provided, comprising:

[0043] The particles may be composed of any material capable of loading and releasing drugs (such as small molecules, therapeutic peptides, proteins, or mRNA) and encapsulated in a gel formed by the polymer building blocks and the crosslinker. The particles may be, for example, silica, liposomes, siRNA complexes, or polymeric materials. The particles can be manufactured using well-known prior art methods, such as emulsion, electrospray, electrostatic complexation, and flow focusing [Abdelaziza, Hadeer M. et al., Journal of Controlled Release 269 374-392 (2018)].

[0044] In some embodiments, the drug is encapsulated in particles comprising a polymeric material selected from the group consisting of polycaprolactone, poly(methacrylic acid), polylactic acid, polyvinylpyrrolidone, poly(lactic-co-glycolic acid) (PLGA), and gelatin. The particles are preferably microparticles and / or nanoparticles, and the particle diameters preferably range from about 10 nm to about 100 μm.

[0045] According to a ninth aspect of the present invention, there is provided a method for producing a protease-responsive drug-loaded hydrogel, comprising the steps of: a) mixing a drug covalently bound to a peptide anchor having a functional group with a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group, thereby covalently bonding the peptide anchor to each functional group of the multi-arm PEG polymer, thereby binding the drug to an arm of the multi-arm PEG polymer; and b) mixing the drug-polymer conjugate of a) with a bifunctional crosslinker comprising a peptide substrate sandwiched between spacer sequences containing functional groups, thereby covalently bonding the functional groups of the polymer building blocks to the functional groups of the bifunctional crosslinker to form a gel; A method of manufacturing is provided, comprising:

[0046] In some embodiments of the ninth aspect of the invention, the method comprises: a) the peptide anchor is protease-cleavable and the cross-linker is not protease-cleavable; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease. It is characterized by:

[0047] In some embodiments, the drug, the particle, the crosslinker, the cleavable anchor, and / or the polymer building block are the same as defined in any aspect of the invention.

[0048] According to a tenth aspect of the present invention, there is provided a method for producing a composite coating material comprising the protease-responsive drug-loaded hydrogel of any one of the aspects of the present invention, comprising the steps of: a) preparing a mixture of a particle-encapsulated drug and a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences comprising functional groups; b) preparing a mixture of a particle-encapsulated drug and a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) mixing the mixture of a) with the mixture of b), adhering the mixture to the coating material, and allowing it to gel; A method of manufacturing is provided, comprising:

[0049] In some embodiments, the dressing is an alginate wound dressing.

[0050] In some embodiments, the method further comprises step d) flash-freezing the composite coating in liquid nitrogen and freeze-drying.

[0051] In some embodiments of the manufacturing method, the drug is a nonsteroidal anti-inflammatory drug (NSAID); the particle comprises poly(lactic-co-glycolic acid) (PLGA); the crosslinker and / or the anchor is cleavable by a protease selected from the group consisting of matrix metalloproteases and serine proteases, and combinations thereof; and the polymer building block comprises a 4-arm or 8-arm PEG-vinylsulfone, a 4-arm or 8-arm PEG-vinylmaleimide, a 4-arm or 8-arm PEG-azide, or a 4-arm or 8-arm PEG-alkyne.

[0052] This generalizable design framework is advantageous because it allows for variation in drug choice and drug loading while maintaining its structural and functional integrity.

[0053] Additionally, this delivery platform is advantageous because it is designed using immunocompatible materials to minimize adverse host responses upon administration in vivo.

[0054] Furthermore, this platform is advantageous because it is versatile enough to accommodate both injectable and topical administration at room temperature. [Brief explanation of the drawings]

[0055] [Figure 1] Figure 1 shows the formation of modular particle-containing hydrogel GEL-iP and the release of drug-loaded particles induced in response to specific protease activity. [Figure 2] This example demonstrates successful gelation of particle-loaded hybrid hydrogels and the induced release of ibuprofen-loaded particles (ibu-PLGA particles) through the action of MMP-9. The addition of a biscysteine ​​peptide as a peptide crosslinker induced gelation (vial A1). In the absence of this crosslinker, gelation did not occur (vial A2). Optical microscopy images (C1) confirmed that gel degradation by MMP-9 activity (vial B1) resulted in the release of drug-loaded particles into the surrounding medium. After 5 days, 200 μL of the hybrid hydrogel was completely degraded. In the absence of MMP-9 activity, the gel remained undegraded (vial B2), and no drug-loaded particles were observed in the surrounding medium (C2). (Scale bar: 50 μm). [Figure 3] This figure shows that the action of MMP-9 induced the release of ibuprofen from the hybrid hydrogel GEL-iP in vitro. When GEL-iP was exposed to MMP-9 (●), the cumulative drug release rate significantly increased compared to the control (■). Addition of an MMP-9 inhibitor (▲) suppressed ibuprofen release. The non-cleavable hybrid hydrogel (scrGEL-iP, ◆) showed a slower release rate. Error bars indicate s.e.m. for n=4 experiments. [Figure 4A-B]This figure shows the effect of drug release from the hybrid hydrogel GEL-iP induced by the action of MMP-9 on macrophage proliferation. A) Schematic diagram of eluate generation, collection, and application to seeded cells. B) Relative metabolic activity of macrophages exposed for 72 hours to eluates collected from the hybrid hydrogels GEL-P, GEL-iP, and scrGEL-iP in the presence or absence of MMP-9 and its inhibitors. Error bars represent s.e.m. for n = 4 experiments. p values ​​were determined by one-way ANOVA with post-hoc Fisher LSD test. (***) indicates p < 0.001, (****) indicates p < 0.0001, and (ns) indicates p > 0.05 (not significant). ibu: ibuprofen; empty PLGA particles: PLGA particles without ibuprofen; ibu-PLGA particles: PLGA particles loaded with ibuprofen; GEL-P: empty PLGA particles embedded in PEG hydrogel cross-linked with cleavable peptide (1) (Figure 14; GCRR-KGPRSLSGK-RRCG; SEQ ID NO: 18); GEL-iP: ibu-PLGA particles embedded in PEG hydrogel cross-linked with cleavable peptide (1) (Figure 14; GCRR-KGPRSLSGK-RRCG; SEQ ID NO: 18); scrGEL-iP: ibu-PLGA particles embedded in PEG hydrogel cross-linked with scrambled peptide (GCRR-KSSRGGPLK-RRCG; SEQ ID NO: 29). [Figure 5A-C]The reactive oxygen species (ROS) activity induced by the hybrid hydrogel GEL-iP and its constituent materials was evaluated in vivo in immunocompetent SKH-1E mice. A) Experimental design: Six material formulations were subcutaneously injected into the back of mice, followed by quantification of ROS activity using bioluminescence imaging. B) Bioluminescence image of a representative mouse on day 3. C) Quantitative ROS activity showing a decrease to background levels by day 5. Error bars indicate sem for n=6 injections. p values ​​were determined by one-way ANOVA with post-hoc Fisher LSD test. (*) indicates p<0.05, and (ns) indicates p>0.05 (not significant). Alginate gel: alginate hydrogel cross-linked with calcium chloride; PEG gel: PEG hydrogel cross-linked with a cleavable peptide; GEL-P: PEG hydrogel cross-linked with a cleavable peptide embedded with PLGA particles; GEL-iP: PEG hydrogel cross-linked with a cleavable peptide embedded with ibu-PLGA particles. [Figure 6A-C] This figure shows that the action of two proteases induced the release of PLGA particles from a multi-protease-cleavable hydrogel. A) Schematic diagram showing the mechanism of the dual-responsive (dual-responsive) combination hydrogel system. B) Photographs (n=3) showing the cleavability of the H2-M2 combination hydrogel over 24 hours: (i) control, (ii) HNE protease-added, (iii) MMP-9 protease-added, and (iv) HNE and MMP-9 proteases-added. C) Measurement of the mean number of PLGA particles released from the combination hydrogel over 24 hours in the absence of proteases or in the presence of one or both proteases (n=3). Error bars represent s.e.m. for n=3 experiments. p values ​​were determined by one-way ANOVA with post-hoc Tukey's test. (***) indicates p<0.001, (ns) indicates p>0.05 (not significant). [Figure 7A-C]Figures show the fabrication of a composite coating incorporating GEL-iP and the in vitro protease-induced release of ibuprofen. A) Schematic diagram of the fabrication of a composite coating using GEL-iP and Cultostat® coating. B) Schematic diagram of MMP-9-induced release of ibuprofen from the composite coating. C) Quantitation of ibuprofen released from the composite coating in response to MMP-9. Error bars represent sem for n=4 experiments. p-values ​​were determined by Student's t-test with Welch's correction. (**) indicates p<0.01. [Figure 8A-C] Figure 1 shows the design of an ibuprofen-conjugated hydrogel with MMP-9 cleavability. A) Schematic of the conjugation reaction and the chemical structure of the ibu-peptide conjugate drawn in ChemDraw®. B) MS spectrum of ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19). C) Schematic of gelation of an ibuprofen-conjugated hydrogel with MMP-9 cleavability and a typical example. [Figure 9A-B] Cleavability of MMP-9-responsive ibuprofen-conjugated PEG hydrogel. A) Degradation of the hydrogel system by MMP-9. B) Schematic and MS spectrum of ibuprofen released by MMP-9. [Figure 10A-B] Figure 1 shows release in response to inflammatory protease stimulation. A) Cumulative amount of free ibuprofen (Libu) increased with increasing MMP-9 concentration. In the absence of MMP-9 protease, no basal release was observed by diffusion (bottom line). B) Specificity of sensitivity of GPRSLSGRRCG (SEQ ID NO: 20) to MMP-9 compared to sensitivity to cathepsin B and HNE. Error bars represent standard error of the mean for n=4 experiments. [Figure 11A-C]This figure demonstrates the tunability of drug loading and drug release rate. The release rate could be tuned by (A) changing the crosslinker (i.e., keeping the anchor H constant, but changing the crosslinker from xH to ​​xM or the control scrambled xM (i.e., xM(scr):GCRR-SSRGGPL-RRCG, SEQ ID NO: 39)) or (B) changing the anchor (i.e., keeping the crosslinker xH constant, but changing the anchor from H to M or the control scrambled M (M(scr):SSRGGPL-RRCG, SEQ ID NO: 40)). C) Changing the number of PEG arms or the weight percent of PEG improved drug loading. The arrow indicates a sudden increase in MMP-9 expression, simulating a sudden relapse. Error bars represent the standard error of the mean for n=4 experiments. The legend is in the format "ibu-anchor_crosslinker." ibu-H_xH refers to ibuprofen-conjugated hydrogels with anchor H and crosslinker xH. [Figure 12A-B] Figure showing mouse models of subcutaneous cavity inflammation with different degrees of severity. A) Schematic diagram of the experimental schedule. Photograph B) and fluorescent image C) show representative mice with three levels of severity. D) Quantitation of fluorescent signals indicating increased MMP activity. E) Quantitation of MMP-9 secretion by ELISA. [Figure 13A-C] Figure showing inflammation-induced drug release in the subcutaneous space of SKH-1E mice. A) Experimental design showing induction of subcutaneous inflammation, subcutaneous injection of drug-conjugated hydrogel into the back of mice, and collection of gel mass 12 hours after hydrogel injection. B) Photographs of representative mouse skin tissues and their gel masses showing three levels of inflammation severity taken on day 3. C) Evaluation of drug release rate. Error bars represent sem for n=8 mice. [Figure 14A-B]This figure shows the results of a qualitative screening of several peptide crosslinkers. Each peptide crosslinker consists of a substrate and two similar spacers, in the form of a spacer-substrate-spacer, as shown in Table 3. A) Gelation was determined when the mixture of PEG-VS, peptide crosslinker, and ibuprofen-loaded particles stopped flowing under gravity and a white hybrid hydrogel formed at the bottom of the vial. B) Relative gelation rates were assessed by comparing the time it took for the liquid mixture to stop flowing. Gelling within 5 minutes was considered "fast," while gelation times exceeding 30 minutes were considered "slow." C) The cleavability of each hybrid hydrogel was confirmed by measuring the amount of drug-loaded particles released into the surrounding medium after exposure to MMP-9 using an optical microscope. (YES) indicates that the number of released particles was significantly higher in the presence of MMP-9, indicating that the hydrogel was cleavable by MMP-9; (NO) indicates that the hydrogel was not cleavable by MMP-9. (ne) indicates a state where no evaluation is being performed. [Figures 15A-E] Scanning electron microscope images of ibuprofen-loaded PLGA particles with different diameters. Different homogenization speeds resulted in particles with average diameters of approximately 46 μm (A), 14 μm (B), 11 μm (C), 6 μm (D), and 4 μm (E). (All scale bars are 20 μm.) [Figure 16] Figure 1 shows the cumulative drug release rate from ibuprofen-loaded PLGA particles of different sizes. As the particle size increased, the drug release rate decreased. All error bars represent sem for n=4 experiments. [Figure 17A-B] Figure 14 shows the in situ formation of PEG gel cross-linked with MMP-9 cleaving peptide (1) (FIG. 14; GCRR-KGPRSLSGK-RRCG; SEQ ID NO: 18). A) Dorsal view of a mouse in which two nodules were formed following subcutaneous injection of a PEG gel precursor solution. B) Photographs of excised skin tissue containing cross-linked PEG gel at two injection sites 15 minutes after injection confirm that PEG gel had formed in situ in the subcutaneous space. [Figure 18A-B] Figure 1 shows the post-injection appearance of a representative mouse. A) Images of a mouse immediately after subcutaneous injection of six different material formulations (alginate gel, PEG gel, ibu-PLGA particles, PLGA particles, GEL-P, and GEL-iP). B) Images of the subcutaneous side of excised skin containing the six material formulations taken 5 days after injection. The PEG gel had disappeared, suggesting its in vivo degradability. [Figure 19] This figure shows the proteolytic functionality of dual-responsive ibuprofen-conjugated PEG hydrogels. Exposure of the gels to MMP-9 and HNE significantly increased the cumulative drug release rate compared to exposure to either MMP-9 or HNE. Furthermore, gels immersed in protease-free buffer exhibited the slowest release rate. Error bars indicate s.e.m. for four experiments.

[0056] For convenience, a list of the references cited herein is provided at the end of the Examples, and the contents of such references are incorporated herein by reference in their entirety. DETAILED DESCRIPTION OF THE INVENTION

[0057] definition For convenience, certain terms employed in the specification, examples, and appended claims are collected here.

[0058] It should be noted that in this specification and the appended claims, the singular forms "a," "an," and "the" include plural references unless the context clearly dictates otherwise.

[0059] Ranges can be expressed herein as from "about" one particular value and / or to "about" another particular value. When such a range is expressed, another embodiment includes the range from the one particular value and / or to the other particular value. Similarly, when values ​​are expressed as approximations, by use of the antecedent "about," it will be understood that the particular value forms another embodiment. It will be further understood that each endpoint of a range has significance both in relation to the other endpoint, and independently of the other endpoint. Also, when more than one value is disclosed herein, it will be understood that for each value, not only that value but also any reference to "about" that particular value is disclosed. For example, when a value is disclosed, "about 10" is also disclosed. It will be understood, as would be appropriate for one of ordinary skill in the art, that when a value is disclosed, "less than or equal to" and "greater than or equal to" that value, as well as possible ranges between those values, are also disclosed. For example, when a value is disclosed, "less than or equal to 10" is disclosed, as well as "greater than or equal to 10." It is also to be understood that each unit between two particular units is also disclosed. For example, if 3 and 10 are disclosed, then 4, 5, 6, 7, 8, and 9 are also disclosed.

[0060] As used herein, the term "amino acid" or "amino acid sequence" refers to an oligopeptide, peptide, polypeptide, protein sequence, or fragment thereof, and refers to a naturally occurring or synthetic molecule. As used herein, when "amino acid sequence" refers to the amino acid sequence of a naturally occurring protein molecule, the term "amino acid sequence" and the like refers not to the complete naturally occurring amino acid sequence associated with the protein molecule being described.

[0061] As used herein, the terms "polypeptide," "peptide," or "protein" refer to one or more chains of amino acids, with the amino acids in each chain covalently linked by peptide bonds. The polypeptide or peptide may comprise multiple chains noncovalently and / or covalently linked by peptide bonds and have the sequence of a native protein, i.e., the sequence of a protein produced by natural cells, particularly non-recombinant cells, or by genetically engineered or recombinant cells. The polypeptide or peptide may also include molecules having the amino acid sequence of a native protein, or molecules having one or more amino acid deletions, additions, and / or substitutions of the native sequence. A "polypeptide," "peptide," or "protein" may comprise a single amino acid chain (referred to as a "monomer") or multiple amino acid chains (referred to as a "multimer").

[0062] As used herein, the term "particle" broadly refers to a material that encapsulates a drug. The particle may be composed of a material capable of carrying and releasing a drug (such as a small molecule, a therapeutic peptide, a protein, or mRNA) and that is encapsulated in a gel formed by the polymer building block and the crosslinker. The particle may be, for example, silica, a liposome, an siRNA complex, or a polymeric material. The particle can be produced using well-known prior art methods, such as emulsion, electrospray, electrostatic complexation, and flow focusing [Abdelaziza, Hadeer M. et al., Journal of Controlled Release 269 374-392 (2018)]. Polymer particles are generally spherical, as shown in Figure 15. Preferred particle sizes for use in the present invention are microparticles and / or nanoparticles with diameters in the nm or μm range. The particle diameter is preferably in the range of 10 nm to 100 μm.

[0063] The term "polymer" or "biopolymer" is defined as a polymerized material of repeating molecular units. The polymer may be a biocompatible polymer selected from the group including polysaccharides (e.g., agarose, dextran), polyphosphazenes, poly(acrylic acid), poly(methacrylic acid), copolymers of acrylic acid and methacrylic acid, poly(alkylene oxidase), poly(vinyl acetate), polyvinylpyrrolidone (PVP), derivatives thereof, and copolymers and polymer blends thereof. When used in drug-loaded polymer particles, the polymer may be selected from the group including polycaprolactone, poly(methacrylic acid), polylactic acid, polyvinylpyrrolidone, poly(lactic-co-glycolic acid) (PLGA), and gelatin. The polymer may be a flexible polymer with mechanical and structural stability suitable for injection, implantation, or implantation (e.g., subcutaneous or subcutaneous implantation). The polymer may or may not be biodegradable. The polymer building blocks of the present invention generally contain multiple arms with functional groups that can interact with the functional groups of a crosslinker to form a gel. Preferred multi-arm building blocks include multi-arm PEG-vinyl sulfone, multi-arm PEG-vinyl maleimide, multi-arm PEG-azide, and multi-arm PEG-alkyne, especially those with four or eight arms.

[0064] As used herein, a "subject" is defined as a vertebrate, particularly a mammal, and more particularly a human. Particularly for research purposes, the subject may be at least one animal model (e.g., a mouse or rat). Particularly for purposes of treatment or prevention of a disease, such as an inflammatory disease, the subject may be a human.

[0065] The term "treatment" as used in the context of the present invention refers to preventative, palliative, therapeutic or curative treatment.

[0066] As used herein, the terms "comprising" or "including" are to be interpreted as specifying the presence of the feature, integer, step, or element described herein denoted by these terms, but not excluding the presence or addition of one or more features, integers, steps, or elements, or groups thereof. In the context of the present disclosure, the terms "comprising" or "including" also encompass the meaning "consisting of." Thus, variations of the term "comprising," such as "comprise" or "comprises," and variations of the term "including," such as "include" or "includes," are similarly broad in meaning.

[0067] While aspects of the present invention will be described in conjunction with the embodiments provided herein, it will be understood that the present invention is not limited to these embodiments. Instead, the present invention encompasses alternatives, modifications, and equivalents to the embodiments described herein, which are within the scope of the present invention as defined by the appended claims. Furthermore, in the following detailed description, specific details are set forth to provide a thorough understanding of the present invention. However, those of ordinary skill in the art will recognize that the present invention can be practiced without these specific details and / or by combining features of individual embodiments. In many instances, detailed descriptions of well-known systems, methods, procedures, and elements are omitted so as not to unnecessarily obscure aspects of the embodiments of the present invention. [Example]

[0068] Those skilled in the art will understand that the present invention can be practiced without undue experimentation by following the methods set forth herein. The methods, techniques, and chemicals set forth herein are as described in the references cited or in protocols found in standard biotechnology or molecular biology textbooks. Standard molecular biology techniques known in the art that are not specifically described herein generally follow the methods described in A Laboratory Manual, Cold Springs Harbor Laboratory, New York (2001).

[0069] Example 1 Materials and methods for fabricating protease-responsive hybrid hydrogels containing drug-encapsulated particles 1.1 Preparation and characterization of PLGA particles Ibuprofen-containing or ibuprofen-free particles were prepared using a water-in-oil emulsion method using 50 / 50 poly(lactic-co-glycolic acid) (PLGA) (Lactel, Pelham, AL) with an intrinsic viscosity of 0.95–1.20 dl / g [Dang, TT et al. Biomaterials 34 (23), 5792–5801 (2013)]. In a typical example, 5 mL of a solution of 40 mg / mL PLGA and 6 mg / mL ibuprofen in dichloromethane was rapidly added to 25 mL of 1% (w / v) polyvinyl alcohol (Sigma-Aldrich, St. Louis, MO, USA) solution and homogenized at various speeds for 60 seconds (L5M-A, Silverson). The resulting suspension was quickly decanted into 75 mL of deionized water, stirred for 60 seconds, and then rotary evaporated for 15 minutes. The suspension was washed three times by centrifugation at 3000 rpm for 30 seconds. The resulting particles were collected, flash-frozen in liquid nitrogen, and lyophilized. The particle size distribution and morphology were examined using a scanning electron microscope (JSM 6390LA, JEOL). 2 mg of particles were dissolved in 1 mL of acetonitrile, and UV absorbance at 240 nm was measured. The measured values ​​were compared with a standard curve of ibuprofen solutions in acetonitrile containing known concentrations of ibuprofen to determine the ibuprofen loading capacity of each particle formulation. The release kinetics from drug-loaded subdomains of ibu-PLGA particles of different sizes were also investigated (Table 1, Figure 15, and Figure 16). For the preparation of GEL-iP, particles with an average diameter of 14 μm and a drug loading of approximately 6 wt% were selected to suppress burst release from the drug-loaded particles.

[0070] [Table 1]

[0071] 1.2 Preparation of hybrid hydrogels containing drug-encapsulated particles We developed a modular drug delivery platform by embedding drug-loaded polymer particles in a protease-cleavable hydrogel (Figure 1). This hydrogel matrix was degraded by proteolytic activity, liberating the embedded particles and ultimately delivering the desired therapeutic payload. The modular nature of this platform allowed us to independently optimize the drug-loading subdomain and the protease-cleavable subdomain, thereby achieving the desired payload release. Specifically, we selected polyethylene glycol (PEG) and poly(lactic-co-glycolic acid) (PLGA) as the main polymer components for the protease-cleavable subdomain and the drug-loading subdomain, respectively. Both of these polymers have been used in clinically approved medical products. PLGA is a synthetic polymer widely used for encapsulating therapeutic agents such as drugs and proteins due to its biodegradability and cytocompatibility [Han, FY et al. Frontiers in pharmacology 7, 185-185 (2016)]. PEG has also been used as a conformal coating for immunoprotection of pancreatic islets [Tomei, AA et al. Proceedings of the National Academy of Sciences 111 (29), 10514 (2014)] and as a component of surgical sealants [Zoia, C. et al. Journal of Applied Biomaterials & Functional Materials 13 (4), 372-375 (2015)]. PEG is a synthetic polymer available in a wide range of molecular weights with diverse multi-arm architectures and functional groups, making it suitable for use in a variety of drug delivery platforms, from systemic and topical to injectable applications [Li, J. and Mooney, DJ Nature Reviews Materials 1, 16071 (2016)].Recently, PEG-based hydrogels have also been demonstrated to be versatile cytocompatible platforms for the stimuli-responsive delivery of multiple drugs and cell-based therapeutics [Badeau, BA Nature Chemistry 10, 251 (2018)].

[0072] Peptide crosslinked hydrogels were prepared by reacting four-arm poly(ethylene glycol)-vinyl sulfone (PEG-VS) (20 kDa, Sigma-Aldrich, St. Louis, MO, USA) with a bis-cysteine ​​peptide (Genscript, Hong Kong) in a stoichiometric ratio. Each precursor was dissolved in triethanolamine (TEOA) buffer (0.3 M) or PBS / NaOH buffer (pH = 10). To prepare 117 μL of a peptide crosslinked hydrogel with a 4.2% (w / v) PEG content, 5 mg of PEG-VS was placed in a glass vial, dissolved in 100 μL of buffer solution, and mixed with 17 μL of the same buffer solution containing the stoichiometric amount of peptide crosslinker. Preliminary screening of the peptide crosslinker was performed using hydrogels with a 4.2% (w / v) PEG content. Subsequent in vitro and in vivo experiments all used hydrogels with a 1.7% (w / v) PEG content. To form hybrid hydrogels in which PLGA particles (with or without ibuprofen) were embedded in peptide-crosslinked hydrogels, the aforementioned precursors were separately dissolved in a buffer solution containing 5% (w / v) PLGA particle suspension. The glass vial containing the liquid mixture of PEG-VS, peptide crosslinker, and PLGA particles was inverted at regular intervals, and gelation was determined when the liquid mixture no longer flowed downward under gravity. Examples of peptide spacer sequences, peptide substrate sequences, and protease susceptibility are shown in Table 2.

[0073] [Table 2]

[0074] 1.3 Screening peptide crosslinkers to achieve optimal gelation and proteolytic cleavage In this study, we designed peptide crosslinkers that could form hydrogels with PEG-VS and retain cleavability by MMP-9 activity. The modular architecture of our hybrid hydrogels allows us to independently design peptide crosslinkers, which are key components of the subdomain that determines MMP-9 cleavability. In a typical example, the desired peptide crosslinker contains four cysteine-containing spacer sequences on either side of an MMP-9-cleavable substrate sequence. This substrate was selected from peptide sequences previously used as the MMP-9-sensitive component of biosensors for MMP-9 detection and as the MMP-9-cleavable linker for chemotherapy drug-loaded nanocarriers [Biela, A. et al. Biosensors and Bioelectronics 68, 660-667 (2015); Samuelson, LE et al. Molecular Pharmaceutics 10 (8), 3164-3174 (2013)]. The thiol group of the cysteine ​​at each terminal spacer is deprotonated to form a thiolate [Friedman, M. et al. Journal of American Chemical Society 87 (16), 3672-3682 (1965)], which reacts with the vinyl sulfone group of PEG-VS via a Michael addition reaction, resulting in gelation [Lutolf, MP and Hubbell., JA Biomacromolecules 4 (3), 713-722 (2003)].

[0075] To identify the optimal peptide crosslinker for GEL-iP, we performed a qualitative screening of eight peptide sequences (Table 3 and Figure 14) to evaluate the effect of substrate and spacer selection on the gelation and cleavability of the hybrid hydrogel. We also investigated the gelation buffer, as the pH of the buffer environment during the gelation reaction may affect the deprotonation of thiols and, ultimately, the crosslinking process. For each combination of substrate, spacer, and buffer, gelation was visually confirmed using the tube inversion method. A stoichiometric amount of peptide crosslinker was added to a glass vial containing a mixture of PEG-VS solution and a suspension of ibu-PLGA particles. A separate vial with the same mixture composition but without the target peptide crosslinker was also prepared as a control. The two vials were inverted at regular intervals, and this procedure was repeated until the mixture in one vial stopped flowing under gravity, indicating gelation. Photographs of glass vials A1 and A2 show typical examples of gelation of hybrid hydrogels (Figure 2). In the presence of the peptide crosslinker, a white hybrid hydrogel formed at the bottom of glass vial A1 and did not flow downward under gravity, confirming crosslinking. This solid white appearance was due to the white ibu-PLGA particles. The precursor mixture in control vial A2, lacking the peptide crosslinker, remained free-flowing and was not considered to have gelled.

[0076] [Table 3]

[0077] 20 μL of each hybrid hydrogel was placed in a 500 μL Eppendorf tube and diluted with PBS buffer (DPBS / modified, calcium-free, magnesium-free, HyClone) containing 3 μg / mL MMP-9 (83 kDa, Merck). TM) was added and incubated at 37°C. As a control experiment, hybrid hydrogels of the same composition were immersed in PBS buffer without MMP-9. After 20 hours of incubation, the medium surrounding the hybrid hydrogels was collected on a coverslip and observed under an optical microscope (Olympus CKX53SF, Japan) to check for the presence or absence of released particles.

[0078] The MMP-9 concentration was set within the range of MMP-9 expression levels found in clinical wound and synovial fluid samples from patients with rheumatoid arthritis and osteoarthritis [Ladwig, GP et al. Wound Repair and Regeneration 10(1) 26-37 (2002); Li, Z. et al. Journal of Diabetes and its Complications 27(4) 380-382 (2013)]. Figure 2 shows typical examples of MMP-9 cleavage. Photographs of vials B1 and B2 in Figure 2 show the appearance of hybrid hydrogels with the same composition as vial A1 after prolonged exposure to MMP-9-containing and MMP-9-free buffers, respectively. In the presence of MMP-9, the white cross-linked hybrid hydrogel seen in vial A1 disappeared, and only a homogeneous, cloudy suspension was observed in vial B1. Optical microscope image C1 confirmed the presence of ibu-PLGA particles in the liquid mixture in vial B1, confirming that the hybrid hydrogel was successfully degraded by MMP-9. Meanwhile, the PBS buffer without MMP-9 added after gelation was observed as a clear liquid phase in vial B2. Optical microscope image C2 also confirmed that no released ibu-PLGA particles were present in vial B2.

[0079] Figure 14 summarizes the results of the qualitative screening of candidate crosslinkers. The screening data in columns (A) and (B) of Figure 14 reveal that the amino acid combination of the peptide crosslinker determines the characteristics of the peptide sequence, which in turn affects the gelation rate. Most of the substrates used in the screening gelled within 5–30 min. Surprisingly, only the substrate AVRWLLTA (SEQ ID NO: 12), a component of peptide (3) (SEQ ID NO: 24) and peptide (8) (SEQ ID NO: 28), did not produce the desired results in terms of the reactivity of the corresponding peptide crosslinker with PEG-VS. Specifically, peptide (3) was gelatinizable by crosslinking with PEG-VS in PBS / NaOH but not in TEOA buffer. We speculate that TEOA, acting as a surfactant, altered the three-dimensional structure or conformation of peptide (3) in aqueous solution [Jones, BH et al. Soft Matter 11 (18), 3572–3580 (2015)], inhibiting gelation. Interestingly, when peptide (8) was dissolved in PBS / NaOH buffer before adding PEG-VS, it self-assembled into a gel, whereas when dissolved in TEOA buffer, it became a milky solution, which is thought to indicate self-assembly of the peptide [Zhou, Q. et al. Progress in Natural Science 19 (11), 1529-1536 (2009)]. This behavior is thought to have inhibited the subsequent reaction of the cysteine ​​thiol of this peptide with the vinyl sulfone of PEG-VS, preventing gelation.

[0080] In addition to the choice of substrate, the design of the spacer (GCRR (SEQ ID NO: 9) or GCRD (SEQ ID NO: 8)) also plays an important role in the crosslinking process. For example, peptide (2) designed with the spacer GCRR (SEQ ID NO: 9) exhibited significantly faster crosslinking rates with PEG-VS than peptide (6) designed with the same substrate but a different spacer, GCRD (SEQ ID NO: 8), in both buffers. Similarly, peptide (4) containing the spacer GCRR (SEQ ID NO: 9) reacted more rapidly with vinyl sulfone in PBS / NaOH than peptide (5) containing the spacer GCRD (SEQ ID NO: 8). This data is consistent with previous reports that the presence of a positive charge (e.g., arginine, R) near the cysteine ​​thiol group increases the crosslinking rate, while the presence of a negative charge (e.g., aspartic acid, D) slows the crosslinking reaction [Lutolf, MP et al. Bioconjugate Chemistry 12 (6), 1051-1056 (2001)]. This is thought to be due to the stabilization of the intermediate thiolate by the former [Roos, G. et al. Antioxidants & Redox Signaling 18 (1), 94-127 (2012)].

[0081] Because environmental pH affects thiol deprotonation and alters the concentration of intermediate thiolates [Lutolf, MP and Hubbell, JA Biomacromolecules 4 (3), 713-722 (2003)], the choice of buffer may also affect the crosslinking rate. TEOA buffer is one of the strongly basic buffers commonly used for Michael additions, but cytotoxicity is a concern. Therefore, we investigated the use of PBS / NaOH buffer as an alternative, more cytocompatible buffer. As shown in Figure 14, all peptides except for the sequence containing the substrate AVRWLLTA (SEQ ID NO: 12) were able to crosslink with PEG-VS in both buffers, although the reaction rates varied. Interestingly, for peptide (5), the crosslinking rate was slower in PBS / NaOH buffer than in TEOA buffer. We speculate that PBS / NaOH, a weakly basic buffer, deprotonates thiols less efficiently than TEOA buffer, which may result in a slower gelation rate. The data obtained on gelation kinetics suggested that even with the same crosslinker design, the appropriate selection of reaction buffer is crucial for successful hydrogel formation.

[0082] Next, column (C) of Figure 14 summarizes the cleavability of each successfully crosslinked hybrid hydrogel in MMP-9 solution. Specifically, hybrid hydrogels crosslinked with peptides (1), (4), (5), and (7) were immersed in an MMP-9-containing buffer, and PLGA particles were released into the surrounding medium, confirming their degradability by MMP-9. Interestingly, hydrogels formed using peptides (2), (3), or (6) were not cleaved by contact with MMP-9, as the number of particles released from the hybrid hydrogels was similarly negligible in both the presence and absence of MMP-9 (not shown). This result was surprising, considering that substrates with these three sequences have been reported as MMP-9-sensitive sites in protease-activated nanocarriers for antitumor drugs [Samuelson, LE et al. Molecular Pharmaceutics 10 (8), 3164-3174 (2013)] and electrochemical impedance sensors [Biela, A. et al. Biosensors and Bioelectronics 68, 660-667 (2015)]. This may be because the spacers flanking the substrate alter the conformation of the peptide sequence, causing steric hindrance and thereby preventing the protease from accessing the substrate cleavage site.

[0083] To design an effective protease-inducible drug delivery platform, the optimal peptide crosslinker must be one that rapidly induces gel formation and maintains protease cleavability. Of the eight peptides screened, three sequences (peptides (1), (4), and (7; Figure 14)) were confirmed to rapidly gel and cleave hydrogels by MMP-9 in all buffers tested. We selected peptide (1), which contains the 17-amino acid sequence GCRR-KGPRSLSGK-RRCG (SEQ ID NO: 18), in combination with PBS / NaOH buffer to prepare the desired hybrid hydrogel, GEL-iP, for further studies.

[0084] 1.4 In vitro drug release kinetics study of hybrid hydrogels containing drug-encapsulated particles In this in vitro release study, two hybrid hydrogels were compared. GEL-iP consisted of ibuprofen-loaded PLGA particles (ibu-PLGA particles) embedded in a hydrogel crosslinked with the cleavable peptide (1) GCRR-KGPRSLSGK-RRCG (SEQ ID NO: 18). ScrGEL-iP consisted of ibu-PLGA particles embedded in a hydrogel crosslinked with the non-cleavable scrambled peptide GCRR-KSSRGGPLK-RRCG (SEQ ID NO: 29). Briefly, 40 μL of GEL-iP was placed in a 1.5 mL tube and immersed in 500 μL of PBS solution containing 3 μg / mL MMP-9 or no MMP-9. 40 μL of scrGEL-iP was exposed to the MMP-9-containing PBS solution alone. A control experiment using GEL-iP was also prepared in which MMP-9 inhibitor I (Merck) was added along with MMP-9. Each tube was placed in a Multi Bio RS-24 rotator (BioSan) set at 37°C and 30 rpm. At designated time points, 10 μL of the liquid mixture was withdrawn from each tube and added to 90 μL of acetonitrile. The resulting samples were passed through a 0.22 μm syringe filter and stored at 4°C. Each tube was replenished with 10 μL of fresh PBS solution containing or without 3 μg / mL MMP-9. After 24 hours, each hybrid hydrogel and the remaining liquid mixture were completely dissolved in acetonitrile. The ibuprofen concentration in all collected samples was quantified by RP-HPLC. The drug release rate at each time point was calculated by normalizing the cumulative amount of drug withdrawn at each time point to the initial amount of drug contained in each tube [Dang, TT et al. Biomaterials 32 (19), 4464-4470 (2011)]. The release rate of each hybrid hydrogel was obtained from the average of four replicate experiments.

[0085] As shown in Figure 3, exposure of GEL-iP to MMP-9 for 4 hours significantly increased the drug release rate to 100%, compared with only 25% in the absence of MMP-9. Furthermore, the co-treatment of MMP-9 with an MMP-9 inhibitor significantly suppressed ibuprofen release. This was due to the selective inhibition of MMP-9 proteolytic activity by this small molecule inhibitor. Therefore, MMP-9 was unable to cleave the peptide crosslinker and degrade the hydrogel matrix, preventing the induced release of ibuprofen. Furthermore, GEL-iP was completely degraded within the first 4 hours in the presence of MMP-9, resulting in 100% ibuprofen release, whereas the hybrid hydrogel crosslinked with a scrambled peptide (scrGEL-iP) only released 40% ibuprofen under the same conditions. In the presence of MMP-9, the amount of ibuprofen released from GEL-iP was greater than that from scrGEL-iP, confirming the function of the cleavable peptide (1) in inducing drug release through MMP-9-mediated degradation of GEL-iP. The data in Figure 3 demonstrate that the induced release of ibuprofen is due to the activity of MMP-9 to cleave its related peptide substrate.

[0086] In vitro evaluation of the inhibitory effect of hybrid hydrogels containing drug-encapsulated particles on macrophages Drugs released from GEL-iP by the action of MMP-9 were evaluated based on their in vitro inhibitory effect on the proliferation of murine macrophages, RAW264.7. Previous studies have revealed that in inflammation-related diseases such as atherosclerosis and obesity-induced adipose tissue inflammation, local macrophage proliferation, rather than monocyte recruitment, is central to focal accumulation of cells [Amano, SU et al. (2014)]. Therefore, macrophage self-renewal is a promising target for therapeutically modulating inflammation.

[0087] Murine macrophages RAW264.7 were cultured in high-glucose DMEM (Gibco Laboratories) containing 10% FBS (Gibco Laboratories) and 1% penicillin / streptomycin (Gibco Laboratories) at 37°C in a 5% CO atmosphere. RAW264.7 macrophages at passages 20-30 were seeded onto 96-well plates (Corning®) at an initial density of 2 x 10 4The cells were seeded at 100 μL per well and incubated at 37°C for 24 hours. Hybrid hydrogels (GEL-iP and scrGEL-iP) made with 1.7% (w / v) PEG and 5% (w / v) ibuprofen-loaded PLGA microparticles were incubated in 500 μL of phenol red-free DMEM (Gibco Laboratories) medium in the presence or absence of 3 μg / mL MMP-9 for 2 hours (Figure 4A). A control hybrid hydrogel, an MMP-9-cleavable hydrogel (GEL-P) containing embedded PLGA microparticles without ibuprofen, was also prepared. In one group of GEL-iP samples, MMP-9 inhibitor I was added along with MMP-9 during incubation. After incubation, 200 μL of culture medium was collected from each hybrid hydrogel formulation as an eluate. This eluate was then added to seeded RAW264.7 macrophages for 72 hours. As a negative control, macrophages were treated with 200 μL of fresh medium. As a positive control, macrophages were treated with 0.6 mg / mL of free ibuprofen solution. The dose of this free ibuprofen solution was normalized to the amount of drug loaded in each hybrid hydrogel. The eluate was then removed from the treated macrophages, and the in vitro metabolic activity of the treated cells was assessed using a WST-1 cell proliferation assay (Abcam) according to the manufacturer's protocol. Specifically, 200 μL of WST-1 reagent (10:1 volume ratio) was added to each well and incubated at 37°C for 3 hours. A 100 μL aliquot of this culture was transferred to a new 96-well plate, and the absorbance at 450 nm and 690 nm was recorded using a microplate reader (SpectraMax M5). To calculate the relative metabolic activity, the absorbance at 450 nm was subtracted from the absorbance at 690 nm (reference wavelength) and the background absorbance to obtain a corrected absorbance value. The relative metabolic activity was calculated using the following formula: Relative metabolic activity = ATest / AControl × 100% where ATest and AControl are the corrected absorbance values ​​of the solutions taken from cells treated with eluate and cells treated with fresh medium, respectively.

[0088] As shown in Figure 4B, the eluate obtained by digesting GEL-iP with MMP-9 completely inhibited macrophage proliferation (approximately 0%), whereas the metabolic activity of cells treated with the eluate obtained from GEL-iP in the absence of MMP-9 was higher (approximately 40%). These data suggest that stimulation of GEL-iP with 3 μg / mL of MMP-9 (which mimics the increased protease expression due to increased inflammation in chronic diseases [Ladwig, G. P. et al. Wound Repair and Regeneration 10(1) 26-37 (2002); Li, Z. et al. (2013)]) increased ibuprofen release and more strongly inhibited macrophage proliferation. Furthermore, the metabolic activity of macrophages treated with the eluate from GEL-iP in the absence of MMP-9 (approximately 40%) was significantly higher than that of macrophages treated with the same amount of free drug solution (simulating uncontrolled systemic drug delivery) (approximately 0%). Thus, in the absence of MMP-9, which mimics non-inflammatory physiological conditions [Roomi, MW et al. (2009)], the amount of ibuprofen released by GEL-iP was less than that contained in the free ibuprofen solution, minimizing the drug's effect on macrophages. This characteristic of GEL-iP may reduce overdosing caused by uncontrolled drug release during systemic drug administration [Youssef, J. et al. Rheumatic Diseases Clinics of North America 42(1) 157-176 (2016)] and potentially alleviate drug-induced side effects in non-inflammatory diseases.

[0089] Furthermore, when an MMP-9 inhibitor was added, macrophage metabolic activity recovered to 20% of the completely inhibited level (approximately 0%) observed in the absence of an MMP-9 inhibitor, demonstrating that active MMP-9 is essential for the desired inhibitory effect of the eluate on macrophages. In the presence of MMP-9, ibuprofen released from GEL-iP completely inhibited macrophage proliferation, whereas macrophages treated with eluate from non-cleavable scrGEL-iP retained 55% of their metabolic activity. Control experiments using an MMP-9 inhibitor and scrGEL-iP indicated that both active MMP-9 and its associated cleavable peptide play an important role in releasing ibuprofen from GEL-iP and regulating macrophage proliferation. Taken together, GEL-iP is a promising drug delivery platform capable of releasing anti-inflammatory agents in response to protease activity and modulating immune cell activity.

[0090] Although our primary goal was to develop a delivery platform that would release drug solely in response to protease activity and minimize release in the absence of this stimulus, some concerns remain regarding the basal release of ibuprofen from GEL-iP in the absence of MMP-9. This is likely due to passive drug diffusion from the surface of ibu-PLGA microparticles, as explained by the partial (approximately 40%) inhibition of macrophages treated with eluate from GEL-iP in the absence of MMP-9 or an MMP-9 inhibitor. However, in clinical settings where anti-inflammatory drug administration is required, some degree of inflammation is often present. Therefore, this basal release of drug may be useful for managing low-level inflammation and associated symptoms such as pain and swelling [Steinmeyer, J. (2000)], thereby minimizing exacerbation of the inflammatory response [Sutherland, ER et al. (2003)]. When inflammation suddenly worsens and MMP-9 activity increases, such as in an arthritis flare-up or a chronic wound infection, GEL-iP releases more ibuprofen, helping to combat the exacerbated inflammation.

[0091] Example 2 In vivo evaluation of protease-responsive hybrid hydrogels containing drug-encapsulating particles 2.1 Management of immunocompetent SKH-1 mice To further evaluate the applicability of GEL-iP as a subcutaneous drug delivery platform, we evaluated the in vivo immunocompatibility of GEL-iP and its constituent materials. Using the immunocompetent mouse model SKH-1E, we investigated the effects of GEL-iP's constituent materials on subcutaneous host responses over a period of up to 5 days (Figure 5). This study was conducted in accordance with an animal experimental protocol (Protocol No. A0343) approved by the Institutional Animal Care and Use Committee (IACUC) of Nanyang Technological University (NTU), Singapore. All animal experiments were conducted in accordance with the National Advisory Committee on Laboratory Animal Research (NACLAR), which complies with the National Institutes of Health's (NIH) Guide for the Care and Use of Laboratory Animals (NIH Publication No. 8023, revised 1978). Breeding mice were purchased from Charles River Laboratories (Wilmington, MA, USA), and 10-week-old female SKH-1E mice (F1) were bred within the NTU facility. Mice were housed in the animal facility of the Lee Kong Chian School of Medicine at NTU under standard conditions with a 12-h light-dark cycle and free access to water and food.

[0092] 2.2 Subcutaneous injection of polymer microparticles Before subcutaneous injection of the materials, mice were anesthetized with inhaled 3% isoflurane in oxygen. Six different material formulations were injected subcutaneously in an array on the dorsal side of each mouse. Specifically, ibu-PLGA particles (50 mg / mL), empty PLGA particles without ibuprofen (50 mg / mL), or 1% (w / v) alginate hydrogel (PRONOVATE) were injected subcutaneously. TM50 μL of PBS buffer containing ibu-PLGA particles (SLG20, FMC BioPolymer) was injected into the mice. For each hydrogel formulation, such as GEL-iP, GEL-P, or PEG hydrogel (PEG gel) crosslinked with peptide (1) (GCRR-KGPRSLSGK-RRCG; SEQ ID NO: 18) without PLGA particles, 50 μL of a solution containing the precursor was injected into the mice. For example, in situ formation of GEL-iP was induced by subcutaneously injecting 50 μL of PBS / NaOH buffer containing PEG-VS, a peptide crosslinker, and ibu-PLGA particles into the back of the mice.

[0093] 2.3 Non-invasive bioluminescence imaging of SKH-1E mice Reactive oxygen species (ROS) activity was quantified using luminol, which is oxidized by ROS and emits a bioluminescent signal, as previously reported [Liu, WF et al. Biomaterials 32 (7), 1796-1801 (2011)]. Briefly, before imaging, mice were intraperitoneally injected with 5 mg of sodium luminol (Sigma Aldrich, St. Louis, MO, USA) dissolved in 100 μL of PBS. Twenty minutes after injection, mice were imaged with an IVIS Spectrum CT system (Caliper Life Sciences) using a 180-second exposure. Regions of interest (ROIs) (cm) around the injection site were then imaged using Living Image 3.1 software. 2 The total radiant flux (photons / sec) of the laser beam was measured.

[0094] In a separate preliminary experiment, the presence of crosslinked hydrogel in the subcutaneous space of excised skin 15 minutes after injection (Figure 17) confirmed in situ gel formation. To evaluate the influence of the GEL-iP components, we also investigated ibuprofen-free PLGA particles and ibuprofen-free PLGA particles. Furthermore, calcium-crosslinked alginate hydrogel (alginate gel), which has previously been reported to be immunocompatible in mice [Liu, WF et al. Biomaterials 32 (7), 1796-1801 (2011)], was used as a negative control.

[0095] Several in vitro and in vivo studies have quantified the activity of reactive oxygen species (ROS) produced by activated phagocytes to characterize the host response to materials [Dang, TT et al. Biomaterials 34 (23), 5792-5801 (2013); Dang, TT et al. Biomaterials 2011, 32 (19), 4464-4470 (2011)]. In this experiment, we used a noninvasive imaging technique to quantify the bioluminescence signal emitted by the oxidation of luminol imaging probes by ROS at the material injection site on days 1, 3, and 5 after material injection (Figure 5A) [Dang, TT et al. Biomaterials 34 (23), 5792-5801 (2013); Dang, TT et al. Biomaterials 32 (19), 4464-4470 (2011); Liu, WF et al. Biomaterials 32 (7), 1796-1801 (2011)]. Figure 5B shows a representative bioluminescence image of a mouse on day 3, and Figure 5C shows the quantitative results of ROS activity induced by various material formulations over a 5-day period. The ROS activity induced by peptide-crosslinked PEG gels without PLGA particles was comparable to that of alginate gels on day 1, confirming the immunoaffinity of the PEG gel in the subcutaneous space. These data confirm that the choice of PBS / NaOH as the buffer for the precursor solution in the preparation of PEG gels does not adversely affect ROS production by immune cells. Furthermore, by day 5, the cross-linked PEG gels in the subcutaneous space of the excised skin had disappeared, suggesting that the PEG gels had completely degraded (Figure 18). Interestingly, the ROS generated by the PLGA-containing formulations was higher than that of the PLGA-free PEG gels and alginate gels in all cases.Although PLGA has been approved by the FDA for use in several drug delivery applications [Han, FY et al. Frontiers in pharmacology 7, 185-185 (2016)], the hydrophobic nature of PLGA may have led to acute inflammation [Seong, S.-Y. and Matzinger, P. (2004)] and the associated increased ROS activity [Dang, TT et al. Biomaterials 34 (23), 5792-5801 (2013)], as observed in this study. However, this PLGA-associated ROS activity was observed on day 1 but eventually decreased to background levels similar to those in control skin by day 5, suggesting that the increase in ROS activity caused by PLGA was transient. Therefore, PLGA is also considered immunocompatible in the subcutaneous space of SKH-1E mice. These findings suggest that the PEG hydrogel and PLGA particles, the components of GEL-iP, are suitable immunocompatible materials for the design of hybrid hydrogels. Furthermore, our selected chemical gelation strategy utilizing bioorthogonal reactions and buffers did not result in an increase in harmful ROS activity.

[0096] Example 3 Multi-responsive protease-cleavable particle-containing hydrogels We demonstrated that multiresponsive protease-cleavable particle-containing hydrogels degraded most rapidly in the presence of multiple disease-specific proteases, demonstrating high targeting specificity for inflammatory diseases. As a typical example, the H2-M2 combination hydrogel, crosslinked with a combination of HNE peptide substrate (H2) and MMP-9 peptide substrate (M2) (Table 4), remained undegraded when exposed to only one protease, i.e., HNE alone or MMP-9 alone, but was completely degraded when both proteases were added (Figure 6A). Furthermore, one-way analysis of variance (ANOVA) of the mean number of PLGA particles in the supernatant showed that the number of particles released from the hydrogels in the presence of two proteases was significantly higher (p<0.001) than from the hydrogels in the presence of one protease or the control condition (Figure 6B). Furthermore, it was shown that the H2 peptide substrate is preferentially cleaved by the HNE protease, and the M2 peptide substrate is preferentially cleaved by the MMP-9 protease.

[0097] [Table 4]

[0098] Example 4 Fabrication of composite coating materials containing protease-responsive hybrid hydrogels containing drug-encapsulated particles 3.1 In vitro drug release kinetics test of composite coating material To demonstrate the versatility of this drug delivery platform for topical application, we incorporated the hybrid hydrogel GEL-iP into Cultostat® wound dressings to fabricate a composite dressing (Figure 7A). This composite dressing ultimately aims to control inflammation and pain by inducing the release of ibuprofen (Figure 7B) through increased MMP-9 levels in chronic wound exudates [Ladwig, G. P. et al. Wound Repair and Regeneration 10(1) 26-37 (2002); Li, Z. et al. Journal of Diabetes and its Complications 27(4) 380-382 (2013)]. To fabricate the composite dressing, peptide (1) (GCRR-KGPRSLSGK-RRCG; SEQ ID NO: 18) and PEG-VS were separately dissolved in a 5% (w / v) suspension of ibu-PLGA particles in PBS / NaOH buffer. These two precursors were mixed, and 20 μL of the precursor mixture was quickly applied to a 6 mm diameter circular sheet of alginate wound dressing Cultostat® (Figure 7A). After gelation, the resulting composite dressing was flash-frozen in liquid nitrogen and lyophilized. The two composite dressings were compared in an in vitro release study. Briefly, a composite dressing containing 20 μL of GEL-iP was placed in a 1.5 mL tube and immersed in 500 μL of a PBS solution containing 3 μg / mL MMP-9, while the other dressing was contacted with a PBS solution without MMP-9.

[0099] The composite coating immediately after formation was saturated with water from the precursor mixture, significantly reducing its ability to absorb further fluid. Therefore, the composite coating was freeze-dried to restore its absorbency. The ibuprofen release capacity of the coating was investigated by immersing it in a buffer solution containing or not containing MMP-9 for 24 hours. After 24 hours of incubation, the composite coating rapidly absorbed the buffer, releasing nearly 100% of the loaded ibuprofen in the presence of MMP-9, whereas the release rate in the absence of MMP-9 was only 56% (Figure 7C). These results demonstrate that the hybrid hydrogel GEL-iP is a versatile triggered drug release platform with potential applications for both injectable (Figure 5) and topical applications (Figure 7).

[0100] Example 5 Modular conjugate hydrogels responsive to single or multiple proteases 5.1 Conjugation of drugs to protease-sensitive peptide anchors We have developed a robust process for the synthesis and purification of new ibuprofen-peptide conjugates. A flow chart of this process is shown in Figure 8A. A brief description of the method is as follows.

[0101] First, ibuprofen, a nonsteroidal anti-inflammatory drug (NSAID), was conjugated to the N-terminus of the peptide sequence GPQGIWGQ-DRCG (SEQ ID NO: 19) by solid-phase peptide synthesis (SPPS) to prepare ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19). Specifically, the Fmoc-protected peptide was manually synthesized on Rink amide resin at a scale of 0.3 mmol / g using standard solid-phase peptide synthesis methods. Prior to coupling the peptide to ibuprofen, the Fmoc protecting group was removed using 20% ​​piperidine. Next, 50 mg of the Fmoc-removed resin was dispersed in 500 μL of DMF along with 9.28 mg of ibuprofen. The reaction was initiated by adding 34.2 μL of a 1 M PyBOP solution in DMF and 6 μL of DIPEA to the resin dispersion. After 18 h, the resin was washed with DMF followed by several washes with dichloromethane (DCM). The resin was then treated with a cleavage cocktail containing 95% trifluoroacetic acid, 2.5% water, and 2.5% triisopropylsilane (TIPS) for 60 minutes at room temperature to cleave ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19). The product was precipitated with cold ether and then dried under vacuum. The identity of this peptide-drug conjugate was confirmed by MS (MS m / z: 731.35 [M+2H] 2+ LC-MS data confirmed that the target drug was successfully conjugated to the N-terminus of the peptide sequence GPQGIWGQ-DRCG (SEQ ID NO: 19), as the observed molecular weight of ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19) was consistent with the theoretically predicted value by ChemDraw® (Figure 8B).

[0102] Second, ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19) was conjugated to the hydrogel using the following procedure. Four-arm poly(ethylene glycol)-maleimide (4-PEG-Mal) (20 kDa, Sigma-Aldrich, St. Louis, MO, USA) was first reacted with ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19) at a 1 / 1 molar ratio. Next, a bis-cysteine ​​peptide (Genscript, Hong Kong) was added to the reaction mixture in a stoichiometric ratio to 4-PEG-Mal (while some of the maleimide groups of 4-PEG-Mal were conjugated to ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19)). Each precursor was dissolved in PBS buffer. As a typical example, to prepare 117 μL of a peptide-crosslinked hydrogel with a PEG content of 4.2% (w / v), 5 mg of 4-PEG-Mal was placed in a glass vial along with 0.40 mg of ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19) and dissolved in 100 μL of PBS. To this solution was then added a stoichiometric amount of the peptide crosslinker dissolved in 17 μL of PBS.

[0103] As shown in the schematic diagram in Figure 8C, adding a bis-cysteine ​​MMP-9-sensitive peptide crosslinker (GCRDGPQGIWGQDRCG; SEQ ID NO: 22) to a buffer solution containing a mixture of 4-arm PEG-maleimide and ibu-GPQGIWGQ-DRCG (SEQ ID NO: 19) resulted in the formation of a crosslinked hydrogel, which did not flow downward under gravity.

[0104] Figure 9 supports the hypothesis that protease-sensitive drug release can be achieved by exposing ibuprofen-conjugated PEG hydrogels to a clinically relevant concentration of proteases, 1 μg / mL. Figure 9A shows that the crosslinked hydrogel was completely degraded after 3 days. LC-MS analysis confirmed the presence of a free ibuprofen-loaded fragment (ibu-GPQG, LC peak at 14.6 min and MS peak at m / z = 546.20) in the supernatant collected after exposing the hydrogel to MMP-9 solution for 2 days (Figure 9B). However, no similar peak was observed in the supernatant collected after exposing the hydrogel to a control buffer without MMP-9. Given that the concentration of active MMP-9 in human wound fluid ranges from 0.3 to 4.8 μg / mL [29-31], the data in Figure 10A support the possibility of drug release at clinically relevant protease concentrations. Furthermore, as shown in FIG. 10B, the peptide sequence GPRSLSGRRCG (SEQ ID NO: 20) as a spacer has higher specificity for MMP-9 than for cathepsin B or human neutrophil elastase (HNE).

[0105] Third, we prepared a hydrogel cleavable by multiple proteases by Michael addition of 4-PEG-Mal with a thiol-containing peptide. Specifically, 5 mg of 4-PEG-Mal was dissolved in 100 μL of PBS buffer, and this solution was reacted with 17 μL of a peptide sequence combination, e.g., 0.005 mg of GRCR-PMAVVQSVP-RCRG (SEQ ID NO: 31) and 0.0045 mg of GRCR-GPRSLSG-RCRG (SEQ ID NO: 32), to prepare 117 μL of 4.2% (w / v) 4-PEG-Mal-containing gel.

[0106] 5.2 Tuning of drug release from conjugate hydrogels in vitro Quantitative data measured by high-performance liquid chromatography (HPLC), shown in Figures 11A and 11B, suggest that drug release from drug-conjugated hydrogels can be tuned by changing the choice of bis-cysteine ​​crosslinker or anchor. Specifically, Figure 11A shows that, when using the same anchor H, changing the crosslinker from peptide xM (GCRR-GPRSLSG-RRCG, SEQ ID NO: 21) to peptide xH (GCRD-GPQGIWGQ-DRCG, SEQ ID NO: 22) results in a decrease in cumulative drug release. Figure 11B also shows that, when using the same crosslinker xH (GCRD-GPQGIWGQ-DRCG, SEQ ID NO: 22), changing the anchor from peptide H (GPQGIWGQ-DRCG, SEQ ID NO: 19) to peptide M (GPRSLSG-RRCG, SEQ ID NO: 20) results in a decrease in cumulative drug release. As shown in Figure 11C, gradually increasing the hydrogel weight ratio from 3 w / v% to 10 w / v% increased the loading, confirming that the total loading could be tuned. It was also found that the 8-arm PEG-maleimide could load more ibuprofen than the 4-arm PEG-maleimide.

[0107] 5.3 Inflammation-induced drug release from conjugate hydrogels in vivo We also obtained preliminary data suggesting that this drug-binding platform can release high doses of drug in response to increasing inflammation in vivo. We created a mouse model of skin inflammation using phorbol 12-myristate 13-acetate (PMA) as a stimulant that can increase MMP expression (Figure 12A). The increase in MMP was detected using the MMPSense probe. 24 h after PMA injection, red areas were observed at the injection site, and the degree of this red area correlated with increasing PMA dose (Figure 12B). Specifically, 4 μg of PMA induced a larger and more pronounced redness than 0.4 μg of PMA, whereas PBS buffer without PMA showed no change in skin appearance. MMP activity was monitored 24 h after PMA injection, and the fluorescent signal was quantified. The MMP activity induced by 0.4 μg PMA was 1.9-fold higher than that induced by 4 μg PMA, indicating that the activity was significantly higher with 4 μg PMA (Figures 12C and 12D). Furthermore, the fluorescent signal at the injection site of PBS buffer without PMA was lowest, confirming the baseline MMP level. Thus, the fluorescent intensity increased with increasing subcutaneous injection doses of PMA. More importantly, this increase in signal intensity correlated with the degree of redness on the dorsal side of the mice. To examine protein expression, skin tissues from the PMA injection site were collected and MMP-9 was quantified using an ELISA kit (Figure 12E). The tendency for increased MMP-9 expression correlated with the severity of inflammation induced by injection of different concentrations of PMA. Specifically, the amount of MMP-9 secretion induced by 4 μg PMA was 5-fold and 10-fold greater than that induced by 0.4 μg or 0 μg PMA, respectively, and the amount of MMP-9 secretion induced by 4 μg PMA was significantly greater. These results indicate that the higher the subcutaneous dose of PMA injected, the more severe the subcutaneous inflammation induced and the greater the amount of MMP-9 secretion.

[0108] To investigate the in vivo inflammation-induced drug release from ibuprofen-conjugated PEG hydrogels, we injected the hydrogel precursor solution into the subcutaneous space of the dorsal inflammatory site of SKH-1E mice to form ibuprofen-conjugated PEG hydrogels in situ (Figure 13A). Drug release was induced by one of three levels of inflammation induced the day before, and the drug release rate was calculated from the amount of ibuprofen remaining in the gel mass 12 h after hydrogel injection. Figure 13B shows that the hydrogels injected without PMA are colorless and transparent, whereas those injected with 4 μg PMA appear yellowish with an unclear border. Figure 13C shows that 60% and 45% of the drug were released from gels exposed to 0.4 μg PMA and PMA-free PBS buffer, respectively, while 70% of the ibuprofen was released from gels exposed to 4 μg PMA. Thus, the amount of ibu-Mf released was positively correlated with the severity of inflammation, confirming a protease-induced release mechanism in vivo. Furthermore, 45% of the drug was released even without PMA. This is likely due to the fact that the injection procedure of PMA and hydrogel administration induced mild inflammation accompanied by increased MMP-9 secretion, resulting in the release of ibu-Mf. Thus, this hydrogel structure demonstrated sensitivity even to mild inflammation. The protease sensitivity of this hydrogel may be tunable by modifying the peptide anchor and / or crosslinker. These results suggest that the drug-conjugated PEG hydrogel can release greater amounts of drug with increasing inflammation severity, potentially addressing differences in the severity of inflammation.

[0109] statistical analysis All statistical analyses and graphs were performed using OriginPro 2017. All comparisons between two experimental groups were performed using Welch's t-test (two-tailed), and comparisons between more than two groups were performed using one-way analysis of variance with Fisher's LSD post-hoc test. P values ​​less than 0.05 were considered significant.

[0110] Example 6 In vitro drug release from modular hybrid hydrogels induced by the action of multiple proteases As a typical example, 20 μL of dual protease-responsive modular conjugate hydrogel was prepared as follows: First, 10 mg of 8-arm PEG-maleimide (8-PEG-MAL, 40 kDa) was placed in a 0.5 mL Eppendorf tube and dissolved in 98.33 μL of PBS buffer solution. 0.17 mg of ibuprofen (Mibu, 1332.7 mmol / mg) conjugated with an MMP-9-sensitive peptide (GPRSLSG-RRCG; SEQ ID NO: 20) was dissolved in 1.67 μL of dimethyl sulfoxide (DMSO), and this solution was mixed with the 8-PEG-MAL solution at a 1:2 stoichiometry. Next, 1.62 mg of an MMP-9 degradable peptide crosslinker (GCRR-GPRSLSG-RRCG; SEQ ID NO: 21) and 1.87 mg of an HNE degradable peptide crosslinker (GRCR-PMAVVQSVP-RCRG; SEQ ID NO: 31) were separately dissolved in 17 μL of PBS buffer and mixed at a 1:1 stoichiometry. The 8-PEG-MAL and Mibu mixture was then reacted with the two peptide crosslinker mixture at a 4:1 stoichiometry to prepare a PEG-peptide hydrogel. The resulting hydrogel solution was homogenized by vortexing and centrifuging for 3 seconds each. The crosslinking time was measured from the start of mixing until no free-flowing solution was observed. The hydrogel was also observed under room light after flicking the tube containing the hydrogel. A hydrogel was considered to have formed if the transparent hydrogel remained attached to the bottom of the tube without disintegrating and no air bubbles were generated inside. This hydrogel was used for all subsequent in vitro gel degradation and drug release experiments.

[0111] Briefly, 20 μL of PEG-peptide hydrogel was placed in a 1.5 mL tube and immersed in 200 μL of PBS containing an enzyme mixture of MMP-9 (2 μg / mL) and HNE (1 μg / mL) or 200 μL of PBS without the enzyme mixture. Separately, 20 μL of PEG-peptide hydrogel was contacted with either MMP-9-containing HEPES solution (2 μg / mL) or HNE-containing HEPES solution (1 μg / mL). Each tube was incubated at 37°C. At designated time points (4, 8, 12, 24, 36, and 48 hours), 5 μL of liquid samples were taken from each tube, placed in 2 mL glass vials, and diluted 4-fold with 15 μL of buffer solution. Each tube was replenished with 5 μL of the corresponding buffer solution. The concentration of digested peptide fragments in all samples was quantified by HPLC. HPLC analysis was performed at room temperature. For HPLC, a 35 / 65 volume mixture of ultrapure water and acetonitrile with 0.1% trifluoroacetic acid was used as the mobile phase at a flow rate of 1.0 mL / min. The detected spectra were converted into drug concentrations, and the cumulative drug release rate was calculated at each time point. The cumulative release rate was calculated by dividing the cumulative drug release amount (the sum of the drug release amount at the measurement time point and the drug disappearance amount up to that point) by the initial amount of peptide used in each hydrogel. The release profile was created by plotting the cumulative drug release rate against the release time. Data for each time point were obtained from the mean and standard deviation of three replicate experiments.

[0112] The release rate of an anti-inflammatory drug from hydrogels in the presence of enzyme solutions containing MMP-9, HNE, or a mixture of MMP-9 and HNE was investigated in vitro. Control experiments were performed using pure buffer solution without enzyme. The cumulative release rates at predetermined time points (4, 8, 12, 24, 36, and 48 hours after the start of exposure to the enzyme solution) were plotted to generate the release profiles shown in Figure 19. As shown in Figure 19, the GEP-peptide hydrogel exposed to a buffer solution containing MMP-9 and HNE for 48 hours exhibited the highest cumulative drug release rate, reaching a maximum of 80%, which was significantly higher than the cumulative release rate (7%) in fresh buffer solution without enzyme. When the hydrogel was exposed to a buffer solution containing only one enzyme, MMP-9, the final cumulative release rate was 71%, which was lower than but close to the maximum cumulative drug release rate observed when MMP-9 and HNE were mixed. When the gel was exposed to protease-free HEPES buffer, the maximum drug release rate after 48 h was 9%, which was lower than the final release rate in the MMP-enzyme mixture solution. This is likely due to the fact that in the non-degradable hydrogel system, resistance to diffusion from the hydrogel prevented sufficient diffusion, resulting in some drug molecules remaining within the hydrogel matrix. A similar problem occurred when the hydrogel was exposed to the HNE enzyme. In this case, the maximum cumulative drug release rate was 24%, which was significantly lower than that in the MMP-9 enzyme solution. This is mainly due to the fact that the HNE enzyme, unlike the MMP-9 enzyme, is unable to cleave the drug conjugate. However, the drug release profile revealed that a small amount of drug molecules were cleaved from the hydrogel and diffused out. This was mainly due to the structural changes of the hydrogel caused by swelling in an aqueous environment, resulting in the cleavage of the peptide crosslinker without hydrogel degradation.

[0113] summary To improve the spatiotemporal control of drug release kinetics in the delivery of anti-inflammatory therapeutics, several strategies have been attempted, including physical encapsulation of drugs or permanent attachment of drugs to polymer backbones [14-17]. However, the release kinetics of previous systems were unable to adapt to changes in the severity of inflammatory disease. Joshi et al. utilized the self-assembly properties of triglycerol monostearate (TG-18) to physically encapsulate corticosteroids within a hydrogel platform and used an enzyme upregulated during arthritis flares as a biological signal to activate drug release

[20] . However, drug release from this platform relies primarily on cleavage of the ester bond in the TG-18 backbone by esterases. Furthermore, the low pH associated with inflammatory conditions [24-26] can lead to nonenzymatic hydrolysis of this ester bond, potentially resulting in undesired nonspecific drug release.

[0114] In this invention, we focused on designing a more efficient drug delivery platform with improved control over the basal release rate and enhanced selectivity and specificity for inflammation-related pathologies, and we were able to demonstrate that the designed platform has the following advantageous properties: (1) A modular system design consisting of multiple embedded subdomains, each with a distinct function, can be individually engineered and replaced to tailor drug loading and release kinetics to specific inflammation-related conditions / diseases by changing the chemical composition of the constituent materials. (2) The basal release rate can be tuned. Specifically, the basal release of drugs can be significantly minimized by covalently binding drugs / modified drugs to inflammation-responsive hydrogels via protease-cleavable peptides, or the basal release can be maintained at a moderate level by using drug-loaded polymer particles as the drug-containing domain. (3) Peptide sequences can be combined to release the loaded substance from the platform upon contact with one or more disease-specific proteases, enhancing the specificity of the platform to allow for the release of a tailored dosage that correlates with the severity of inflammation in the disease. These advantages are confirmed by the following examples.

[0115] Modular microparticle-containing hydrogels responsive to single or multiple proteases We developed a modular hybrid hydrogel that can induce the release of anti-inflammatory drugs in response to increased protease activity associated with inflammatory diseases. The hydrogel matrix was proteolytically degraded by protease activity, liberating embedded particles and ultimately delivering the desired therapeutic payload. Because of the modular nature of this hybrid hydrogel, the protease-cleavable subdomain and drug-loading subdomain could be independently optimized to promote hydrogel formation and matrix metalloproteinase-9 (MMP-9) cleavage, ultimately delivering the desired payload at tunable release rates. In vitro studies confirmed that protease-mediated drug release from the hybrid hydrogel system effectively suppressed TNF-α production by inflammatory macrophages, suggesting the potential for mitigating drug-induced cytotoxicity. We also used noninvasive imaging to monitor reactive oxygen species activity during biomaterial-induced host responses and confirmed that the hybrid hydrogel and its constituent materials did not induce adverse immune responses 5 days after subcutaneous administration in immunocompetent mice. We then incorporated this hybrid hydrogel into a commercially available wound dressing and confirmed that drug release occurred upon contact with MMP-9. These results suggest that this hybrid hydrogel is a versatile platform for on-demand drug delivery via injection or topical application to modulate inflammation in chronic diseases.

[0116] Modular conjugate hydrogels responsive to single or multiple proteases We also developed a modular hydrogel system conjugated with an anti-inflammatory drug. We confirmed that the release of the therapeutic drug was induced by stimulation with one or two types of proteases in this hydrogel system. In some embodiments, the drug loading of the drug-conjugated hydrogel system could be increased by modifying the composition of the polyethylene glycol backbone of the hydrogel. Furthermore, the drug release rate could be tuned by changing the protease-cleavable anchor and crosslinker. Furthermore, we confirmed in vivo protease-induced drug release using chemically induced subcutaneous inflammation models with different severity levels.

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[0118] The present invention includes the following inventions. [1] A protease-responsive drug-loaded hydrogel, a) Drugs encapsulated in particles; b) a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences containing functional groups; Including, A protease-responsive drug-loaded hydrogel, characterized in that the polymer building block (b) forms a gel in the presence of the protease-cleavable crosslinker (c), and the particles (a) are encapsulated in the gel. [2] a) at least one second bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by spacer sequences comprising functional groups; and / or b) at least one bifunctional protease-resistant crosslinker containing a protease-resistant substrate; Further including, The protease-responsive drug-loaded hydrogel according to [1] above, characterized in that the protease-cleavable substrate a) is sensitive to a protease different from the protease to which the crosslinker c) is sensitive. [3] The protease-responsive drug-loaded hydrogel according to [1] or [2], wherein the drug is encapsulated in particles comprising a material selected from the group consisting of silica, liposomes, siRNA complexes, and polymeric materials such as polycaprolactone, poly(methacrylic acid), polylactic acid, polyvinylpyrrolidone, poly(lactic-co-glycolic acid) (PLGA), and gelatin. [4] The protease-responsive drug-loaded hydrogel according to any one of [1] to [3] above, wherein the polymer building block comprises multi-arm PEG-vinyl sulfone, multi-arm PEG-maleimide, multi-arm PEG-azide, or multi-arm PEG-alkyne. [5] A protease-responsive drug-loaded hydrogel, a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, A protease-responsive drug-loaded hydrogel, characterized in that the drug and the arms of the multi-arm PEG polymer are covalently bonded via the functional groups of the peptide anchor, and a gel is formed by covalently bonding the functional groups of the polymer building block and the functional groups of the bifunctional crosslinker. [6] a) the crosslinker is not cleavable by proteases; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease. The protease-responsive drug-loaded hydrogel according to [5] above, characterized by: [7] The protease-responsive drug-loaded hydrogel according to [5] or [6], wherein the polymer building block comprises multi-arm PEG-vinyl sulfone, multi-arm PEG-vinyl maleimide, multi-arm PEG-azide, or multi-arm PEG-alkyne. [8] The protease-responsive drug-loaded hydrogel according to [7] above, which contains 2 to 12 wt % of multi-arm PEG-vinylmaleimide. [9] The protease-responsive drug-loaded hydrogel according to any one of [1] to [8] above, wherein the multi-arm PEG polymer has 3 to 8 arms.

[10] The protease-responsive drug-loaded hydrogel according to any one of [1] to [9] above, wherein the drug is an anti-inflammatory agent.

[11] The protease-responsive drug-loaded hydrogel according to

[10] above, wherein the drug is a non-steroidal anti-inflammatory drug (NSAID).

[12] The protease-responsive drug-loaded hydrogel according to

[10] or

[11] , wherein the protease is a protease whose expression level increases during inflammation and is selected from the group including matrix metalloproteases, serine proteases, cysteine ​​proteases, and aspartic acid proteases.

[13] The protease-responsive drug-loaded hydrogel according to any one of [1] to

[12] above, wherein the spacer sequences sandwiching the substrate comprise at least one cysteine ​​residue, at least one lysine residue, and / or at least one azide-containing unnatural amino acid or alkyne-containing unnatural amino acid.

[14] The protease-responsive drug-loaded hydrogel according to

[13] , wherein the spacer sequences sandwiching the substrate comprise a sequence consisting of 1 to 6 amino acids.

[15] The protease-responsive drug-loaded hydrogel according to any one of [1] to

[14] , wherein the protease-cleavable substrate is sensitive to a protease selected from the group consisting of matrix metalloproteinases such as metalloproteinase-9 (MMP-9), MMP-2, MMP-7, and MMP-12, cathepsins such as cathepsin K, cathepsin B, and cathepsin S, human neutrophil elastase (HNE), caspases, and urokinase.

[16] The protease-responsive drug-loaded hydrogel according to

[15] , wherein the protease-cleavable substrate is selected from the group consisting of an MMP-9 substrate comprising the amino acid sequence represented by KGPRSLSGK (SEQ ID NO: 30), GPRSLSG (SEQ ID NO: 10), LGRMGLPGK (SEQ ID NO: 11), AVRWLLTA (SEQ ID NO: 12), or GPQGIWGQ (SEQ ID NO: 13), an HNE substrate comprising APEEIMDRQ (SEQ ID NO: 14) or PMAVVQSVP (SEQ ID NO: 15), a cathepsin B substrate comprising GRRGLG (SEQ ID NO: 16) or DGFLGDD (SEQ ID NO: 17), and combinations thereof.

[17] A composition comprising the protease-responsive drug-loaded hydrogel according to any one of [1] to

[16] above, formulated for injection or topical administration.

[18] A dressing material comprising the protease-responsive drug-carrying hydrogel according to any one of [1] to

[16] above.

[19] Use of the protease-responsive drug-loaded hydrogel according to any one of [1] to

[16] above or the composition according to

[17] above as an injection or external dressing for treating a subject in need thereof.

[20] A method for treatment, comprising administering an effective amount of the protease-responsive drug-loaded hydrogel described in any one of [1] to

[16] above or the composition described in

[17] above to a subject in need of such treatment.

[21] a) Drugs encapsulated in particles; b) a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences containing functional groups; Including, a kit, wherein a) to c) are as defined in any of the preceding items; or a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm PEG polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, A kit characterized in that a) to c) are the same as those defined in any of the preceding items.

[22] A method for producing a protease-responsive drug-loaded hydrogel, comprising: a) mixing a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG) with drug-loaded particles; b) mixing the drug-loaded particles with a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by spacer sequences comprising functional groups; and c) mixing the mixture of a) with the mixture of b) so that the polymer building blocks of a) form a gel in the presence of the protease-cleavable crosslinker of b), and the drug-loaded particles are encapsulated in the gel. A manufacturing method comprising:

[23] A method for producing a protease-responsive drug-loaded hydrogel, comprising: a) mixing a drug covalently bound to a peptide anchor having a functional group with a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group, thereby covalently bonding the peptide anchor to each functional group of the multi-arm PEG polymer, thereby binding the drug to an arm of the multi-arm PEG polymer; and b) mixing the drug-polymer conjugate of a) with a bifunctional crosslinker comprising a peptide substrate sandwiched between spacer sequences containing functional groups, thereby covalently bonding the functional groups of the polymer building blocks to the functional groups of the bifunctional crosslinker to form a gel; A manufacturing method comprising:

[24] a) the peptide anchor is protease-cleavable and the cross-linker is not protease-cleavable; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease. The method according to

[23] above, characterized in that:

[25] The method according to any one of

[22] to

[24] , wherein the drug, the polymer particle, the crosslinker, the cleavable anchor, and / or the polymer building block are the same as those defined in any one of [1] to

[16] .

[26] A method for producing a composite dressing material containing the protease-responsive drug-carrying hydrogel according to any one of [1] to

[16] , a) preparing a mixture of a particle-encapsulated drug and a bifunctional protease-sensitive crosslinker comprising a protease-cleavable substrate flanked by two spacer sequences comprising functional groups; b) preparing a mixture of a particle-encapsulated drug and a polymer building block comprising a functionalized multi-arm polyethylene glycol (PEG); and c) mixing the mixture of a) with the mixture of b), adhering the mixture to the coating material, and allowing it to gel; A manufacturing method comprising:

[27] The method according to

[26] , wherein the dressing is an alginate wound dressing.

[28] The method according to

[26] or

[27] , further comprising step d) of flash-freezing the composite coating in liquid nitrogen and freeze-drying.

[29] The method according to any one of

[22] to

[28] above, wherein the drug is a nonsteroidal anti-inflammatory drug (NSAID); the particle comprises poly(lactic-co-glycolic acid) (PLGA); the crosslinker and / or the anchor is cleavable by a protease selected from the group consisting of matrix metalloproteinases and serine proteases, and combinations thereof; and the polymer building block comprises a 4-arm or 8-arm PEG-vinylsulfone, a 4-arm or 8-arm PEG-vinylmaleimide, a 4-arm or 8-arm PEG-azide, or a 4-arm or 8-arm PEG-alkyne.

Claims

1. A protease-responsive drug-loaded hydrogel, comprising: a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, A protease-responsive drug-loaded hydrogel, characterized in that the drug and the arms of the multi-arm PEG polymer are covalently bonded via the functional groups of the peptide anchor, and a gel is formed by covalently bonding the functional groups of the polymer building block and the functional groups of the bifunctional crosslinker.

2. a) the cross-linker is not protease-cleavable; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease. The protease-responsive drug-loaded hydrogel according to claim 1, characterized in that:

3. 3. The protease-responsive drug-loaded hydrogel of claim 1, wherein the polymer building block comprises multi-arm PEG-vinyl sulfone, multi-arm PEG-vinyl maleimide, multi-arm PEG-azide, or multi-arm PEG-alkyne.

4. The protease-responsive drug-loaded hydrogel of claim 3, comprising 2 to 12 wt% multi-arm PEG-vinylmaleimide.

5. The protease-responsive drug-loaded hydrogel according to any one of claims 1 to 4, wherein the multi-arm PEG polymer has 3 to 8 arms.

6. The protease-responsive drug-loaded hydrogel according to any one of claims 1 to 5, wherein the drug is an anti-inflammatory agent.

7. The protease-responsive drug-loaded hydrogel of claim 6, wherein the drug is a non-steroidal anti-inflammatory drug (NSAID).

8. The protease-responsive drug-loaded hydrogel of claim 6 or 7, wherein the protease is a protease whose expression increases during inflammation and is selected from the group including matrix metalloproteases, serine proteases, cysteine ​​proteases, and aspartic acid proteases.

9. The protease-responsive drug-loaded hydrogel of any one of claims 1 to 8, wherein the spacer sequences flanking the substrate comprise at least one cysteine ​​residue, at least one lysine residue, and / or at least one azide-containing unnatural amino acid or alkyne-containing unnatural amino acid.

10. The protease-responsive drug-loaded hydrogel of claim 9, wherein the spacer sequences flanking the substrate comprise a sequence consisting of 1 to 6 amino acids.

11. The protease-responsive drug-loaded hydrogel according to any one of claims 1 to 10, wherein the protease-cleavable substrate is sensitive to a protease selected from the group consisting of matrix metalloproteases such as metalloprotease-9 (MMP-9), MMP-2, MMP-7, and MMP-12, cathepsins such as cathepsin K, cathepsin B, and cathepsin S, human neutrophil elastase (HNE), caspases, and urokinase.

12. The protease-responsive drug-loaded hydrogel of claim 11, wherein the protease-cleavable substrate is selected from the group consisting of an MMP-9 substrate having an amino acid sequence represented by KGPRSLSGK (sequence number 30), GPRSLSG (sequence number 10), LGRMGLPGK (sequence number 11), AVRWLLTA (sequence number 12), or GPQGIWGQ (sequence number 13), an HNE substrate having an APEEIMDRQ (sequence number 14) or PMAVVQSVP (sequence number 15), a cathepsin B substrate having GRRGLG (sequence number 16) or DGFLGDD (sequence number 17), and combinations thereof.

13. A composition comprising the protease-responsive drug-loaded hydrogel of any one of claims 1 to 12, formulated for injection or topical administration.

14. A dressing material comprising the protease-responsive drug-loaded hydrogel according to any one of claims 1 to 12.

15. Use of the protease-responsive drug-loaded hydrogel of any one of claims 1 to 12 or the composition of claim 13 in the manufacture of a medicament for treating a subject in need thereof.

16. The use according to claim 15, wherein the pharmaceutical is an injection or an external dressing.

17. a) a drug covalently attached to a functionalized protease-cleavable peptide anchor; b) a polymer building block comprising a multi-arm PEG polymer having at least one functional group; and c) a bifunctional crosslinker comprising a peptide substrate flanked by spacer sequences containing functional groups; Including, A kit characterized in that a) to c) are as defined in any preceding claim.

18. A method for producing a protease-responsive drug-loaded hydrogel, comprising: a) mixing a drug covalently bound to a peptide anchor having a functional group with a polymer building block comprising a multi-arm polyethylene glycol (PEG) polymer having at least one functional group, thereby covalently bonding the peptide anchor to each functional group of the multi-arm PEG polymer, thereby binding the drug to an arm of the multi-arm PEG polymer; and b) mixing the drug-polymer conjugate of a) with a bifunctional crosslinker comprising a peptide substrate sandwiched between spacer sequences containing functional groups, thereby covalently bonding the functional groups of the polymer building blocks to the functional groups of the bifunctional crosslinker to form a gel; A manufacturing method comprising:

19. a) the peptide anchor is protease-cleavable and the cross-linker is not protease-cleavable; or b) the peptide anchor is cleavable by a protease, and the cross-linking agent is cleavable by a protease that is the same as or different from the protease that cleaves the peptide anchor; and / or c) The drug-loaded hydrogel contains multiple types of crosslinking agents, at least one of which is cleavable by a different protease.

20. The method of claim 18, wherein:

20. 20. The method according to claim 18 or 19, characterized in that the drug, the polymer particle, the crosslinker, the cleavable anchor and / or the polymer building block are the same as defined in any one of claims 1 to 12.

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