Perforation is a hole that develops through the wall of a body organ. A gastric perforation (GP) is a full-thickness injury of the stomach wall with spillage of the gastric contents into the general peritoneal cavity ^[1][51]. Severe complications are often associated with GP including bleeding, sepsis, chemical peritonitis, bowel infarction, wound infection, and multi-organ failure ^[2][3][4][52,53,54]. The most common cause of gastric perforation is peptic ulcer disease, a chronic disease that results from an imbalance between endogenous protective factors of the gastric mucosa and aggressive factors, with a lifetime prevalence of 5–10% in the general population ^[5][55]. GP also arises from trauma, malignancy, intrinsic gastric pathology, or endoscopy-related interventional procedures including EMR and ESD ^[1][51]. An intestinal perforation (IP) is a loss of continuity of the bowel wall, resulting from a variety of disease processes such as ischemia, infection, erosion, and physical disruption, and causing complications such as sepsis, peritonitis, and anastomotic leakage ^[6][56].

Non-surgical management of perforation usually involves analgesia, intravenous antibiotics, and proton pump inhibitory medications (PPIs), but faces the risk of higher mortality rate if it fails ^[5][55]. Therefore, current treatment primarily relies on endoscopic closure and laparotomy for perforations less than 10 mm and open surgery for larger perforations ^[5][7][55,57]. Direct closure of the perforation with interrupted sutures and an omental pedicle plug is the most commonly used technique. However, this procedure results in a 7% suture leak rate with laparoscopic repair and is usually associated with severe tissue damage and inflammation caused by deep piercing and ischemia ^[4][5][54,55]. As a result, the estimated postoperative mortality rate for perforation ranges from 1.3% to 20%, with the 90-day mortality rate reaching up to 30% ^[5][55]. Therefore, there is an imperative need for the development of new biomaterials and techniques to improve perforation sealing and facilitate sutureless GI repair.

To form fluid-tight sealing of GI perforation, an ideal hydrogel adhesive requires: (1) extremely strong and stable long-term wet adhesion; (2) mechanically robust network; (3) tolerance for extreme pH conditions (especially for gastric perforation); (4) excellent interfacial toughness, ductility, and fatigue resistance; (5) ease of application and use; and (6) biocompatibility. Unlike the hydrogels designed for GI hemostasis and wound healing, hydrogels here focus more on enhancing the mechanical properties (G′ > 6 kPa, interfacial toughness > 50 J/m², burst pressure > 120 mmHg) and strengthening wet adhesion (adhesive strength > 10 kPa for over 7 days), as they need to sustain fluid pressure and prevent leakage before the perforation completely heals. Strong and durable interfacial bonds need to be established between hydrogels and the tissue surface to prevent adhesive failure, whereas robust, tough, and acid-tolerating networks need to be constructed for the bulk hydrogel to prevent cohesive failure. Additionally, the adhesive may be delivered endoscopically or laparoscopically as a glue or through open surgery as a patch depending on the location and size of the perforation.

Hydrogen bonds can endow materials with high toughness, elasticity, and self-healing properties owing to their dynamic nature. By incorporating substantial amounts of free hydrogen bonding groups, such as carboxylic acid, amine, and alcohol, into the network, strong adhesion interfaces form between the hydrogel and the tissue surface. A hydrogel adhesive prepared by free radical polymerization of a bi-carboxyl-containing vinyl monomer N-acryloyl aspartic acid (AASP) shows good toughness, robust elasticity, fatigue resistance, strong adhesion to various tissues (120 kPa), and stability in simulated gastric fluid for more than 7 days ^[8][58]. The rich carboxyl groups on the side chains ensure both the formation of the hydrogel network and the interaction with the polar groups on the tissue interface ^[8][58]. Physically crosslinked polyelectrolytes also form strong hydrogen bonding interactions by diffusing into the tissue. A self-gelling and adhesive polyethyleneimine (PEI) and polyacrylic acid (PAA) powder effectively seals gastric perforation within 2 s in a rat model ^[9][59]. Due to the water-induced gelation mechanism, the PEI/PAA powder absorbs interfacial water to enhance wet adhesion and can be delivered endoscopically to fit irregularly shaped target sites. However, the hydrogel does not survive extreme pH environments (pH < 2 or pH > 11) because of its high charge density, limiting its application to nondigestive periods and only the lower GI tract ^[9][59].

Additionally, hydrogels based on dynamic covalent bonds are attractive due to their inherent self-healing ability and high cohesion on account of the strong covalent bonds ^[10][14]. A hydrogel patch adhesive, prepared through Schiff base reaction between tetra-PEG-CHO and carboxymethyl chitosan (CMCS), displays high storage modulus (25 kPa), high burst pressure (217.5 mmHg), and rapid self-healing under physiological conditions ^[11][38]. Wang et al. report an injectable hydrogel assembled from an ABA triblock copolymer composed of a middle PEG block and terminal temperature-responsive poly(N-isopropylacrylamide-co-N-acryloyl-6-aminocaproic acid) P(NIPAM-co-NA6ACA) blocks ^[12][39]. Upon application, the polymer solution transforms into a physical supramolecular hydrogel due to the hydrophobic interactions between the collapsed P(NIPAM-co-NA6ACA) blocks. The hydrogel self-heals in the acidic environment due to the synergy of hydrogen bonding and hydrophobic interactions, bestowing an added advantage. Its adhesive strength to porcine stomach, however, is relatively weak (6 kPa) owing to its mere hydrogen bonding adhesive mechanism ^[12][39].

As discussed above, biomimetic catechol-based mussel-inspired hydrogels are extensively studied for wet adhesion. However, it is a challenge to fabricate catechol-based hydrogels for perforation sealing because strong adhesion and cohesion of the hydrogel require drastically different pH conditions, respectively ^[13][14][60,61]. Alternatively, tannic acid (TA) is a natural plant-based polyphenol containing a high density of pyrogallol and catechol groups that functions as a physical crosslinker as well as an adhesive moiety to facilitate both cohesion and adhesion in acidic conditions ^[14][15][61,62]. Its dendritic structure provides multiple bonding sites for hydrogen bonding, ionic bonding, and hydrophobic interactions ^[14][61]. Additionally, TA adhesion increases when exposed to oxidants ^[13][60]. A gelatin methacrylate (GelMA) containing TA exhibits high stiffness (4.6 MPa), high adhesive strength (80 kPa), large deformability (277% elongation), and self-healing property ^[15][62]. Its stability in acidic conditions, however, is questionable due to its hydrolyzable ester linkages in GelMA. Acid-tolerant hydrogel adhesives based on dopamine-modified poly(γ-glutamic acid) (γ-PGA) and TA also exhibit high storage modulus (5 kPa), high adhesive strength (50 kPa), and large deformability (800% elongation). Notably, due to its abundant phenolic hydroxyl groups and complete physical crosslinking mechanism, the hydrogels display stable and robust adhesion in simulated gastric juice (pH = 1) for 7 days, significantly surpassing fibrin glue in acid-tolerating capacities ^[13][60].

To further enhance wet adhesion, interfacial water dehydration is necessary to remove the gap between the adhesive and the tissue ^[16][63]. There exist two distinct strategies: (1) increasing polymer hydrophobicity to break the hydration layer and displace the water at the molecular level; and (2) utilizing polymer hydrophilicity to absorb interfacial water at the microscopic level ^[16][63]. Glue-type injectable hydrogels usually adopt the first strategy to enhance molecular-level interfacial bonding as the polymers diffuse across the interfacial water and form mechanical interlocking with the tissue layer. Hydrophobically modified Alaska pollock gelatin microparticle-based wound dressings are of interest for treating not only acute GI bleeding but also ESD-induced perforation ^[17][18][28,44]. Gradual fusion of the microparticles occurs during the first 60 min of application via hydrophobic interactions to increase the burst strength of the hydrogel and improve its underwater stability on duodenum tissue ex vivo ^[17][28]. In another study, silica nanoparticles coat hydrogels to enhance the adsorption and entanglement between the hydrogel and the tissue surface ^[2][52]. The wet adhesive capability of biomimetic TA-based hydrogels is further strengthened by the inclusion of hydrophobic moieties. A self-hydrophobized adhesive co-assembled from disulfide-bond-hydrolyzed silk sericin protein and TA achieves extremely robust (>100 kPa for tissues) and durable (>7 days) underwater adhesion and seals a mouse small intestine perforation instantly ^[19][45]. The dissociation of disulfide bonds in silk sericin protein induces hydrophobic amino acid eversion, thereby leading to the self-aggregation of hydrophobic chains once exposed to water, repelling interfacial water, and enhancing subsequent interfacial physical crosslinking effects ^[19][45].

Patch-type hydrogels, on the other hand, frequently adopt the second strategy, also known as the dry-crosslinking mechanism, to remove interfacial water and facilitate subsequent crosslinking. Upon contact with wet surfaces, a dry double-sided tape made from the combination of a biopolymer (gelatin or chitosan) and crosslinked PAA grafted with NHS ester instantly dries the interfacial water and swells ^[20][40]. Temporary adhesion via hydrogen bonding occurs within the first 5 s, followed by covalent adhesion between the NHS ester groups and the primary amine groups on the tissue surface, enabling its potential application to seal air-tight lung lobes as well as fluid-tight perforated stomachs and small intestines ^[20][40]. Additionally, the degradation and mechanical properties of the network are controlled by tuning the composition of the biopolymer component ^[20][40]. To integrate both strategies of wet adhesion, Liu et al. describe a half-dry adhesive for rapid gastric perforation and traumatic pneumothorax sealing ^[7][57]. The hydrogel first repels the majority of interfacial liquid based on its moderate hydrophilicity, then absorbs trapped residues and vicinal tissue fluid to enhance topological adhesion, and finally bonds to the tissue surface through physical interactions ^[7][57]. With the combined strategy, the PAA-silk fibroin (SF) based adhesive achieves excellent adhesion energy (600 J/m²) and burst pressure (1500 mmHg) for more than 24 h.

Interpenetrating polymer network (IPN) hydrogels are polymer composites composed of two or more crosslinked networks that are topologically entangled and cannot be separated without disrupting existing chemical bonds ^[21][64]. Compared to single network hydrogels, IPNs possess more widely controllable physical properties and more versatile functionalities ^[22][65]. Due to the multitude of design requirements needed for hydrogels to target GI perforation, various IPN hydrogels are being developed to improve on at least one property of the single network, such as mechanical reinforcement, adhesion enhancement, or better energy dissipation. These IPN hydrogels typically comprise a covalently crosslinked rigid network and a physically crosslinked soft network. The covalent network acts to reinforce the mechanical properties, stabilize the polymer network, and prevent dissociation in extreme pH environments ^[20][23][40,66]. The physical network allows for energy dissipation, good ductility and flexibility, and self-healing properties ^[4][54].

Sodium alginate (SA) physically crosslinked with calcium ions (Ca²⁺) serves as a second network in a poly(acrylamide) (PAM)-based covalent network to enhance energy dissipation in multiple instances ^[2][24][52,67]. One of the challenges for GI perforation repair lies in the excessive swelling of hydrogels, which may cause a mismatch strain between the hydrogel and the tissue, thereby reducing the overall adhesion performance ^[24][67]. To address this issue, nano-hydroxyapatites are embedded into the PAM-based network to function as ionic nano-reservoirs and gradually release Ca²⁺ in acidic environments, building a second SA-based network to inhibit swelling of the hydrogel in gastric juice ^[24][67]. Silk fibroin (SF) is a natural multi-domain protein that is also being incorporated into IPN hydrogel adhesives owing to its superior strength and stretchability ^[25][68]. Liu et al. report a half-dry adhesive for wet adhesion consisting of a SF semi-interpenetrating network and a PAA covalent network ^[7][57]. After the addition of acrylic acid, the heavy chains of SF can rearrange from α-helical conformations to antipolar-antiparallel beta-sheets (β-sheets), providing additional physical crosslinks and enhancing the bulk toughness of the hydrogel ^[7][57]. Dynamic covalent bonding networks also complement covalent networks to enhance mechanical strength and provide energy dissipation due to their thermodynamically controlled reversible junctions ^[10][14]. Chen et al. report an injectable hydrogel consisting of a bioactive, transglutaminase (TG)-crosslinked gelatin network and a dynamic, borate-crosslinked poly-N-[tris(hydroxymethyl)methyl]acrylamide (PTH) network ^[4][54]. The primary gelatin network covalently crosslinks with tissue amines via TG to provide adhesion whereas the secondary PTH network favors energy dissipation through its reversible boronic-ester bonds ^[4][54].

Due to the versatility and flexible functionalities of IPN hydrogels, researchers are adopting various strategies to construct IPN hydrogels with not only enhanced mechanical properties but also acid-tolerating capabilities for GI perforation applications. Bian et al. describe a fit-to-shape sealant enhanced by photo-initiated crosslinking to treat wounds inside the stomach with the existence of gastric acid ^[23][66]. The primary network of the hydrogel is based on dynamic Schiff-base linkages between chitosan and benzaldehyde-terminated PEG to endow the hydrogel with shear-thinning injectable properties. The secondary network forms via photopolymerization among the vinyl groups on maleic-modified chitosan, polyethylene glycol diacrylate (PEGDA), and dopamine methacrylamide (DMA) to enhance mechanical properties and provide wet adhesion through the catechol moieties ^[23][66]. The resulting hydrogel maintains a high adhesive strength (35 kPa) and an integral gel state in extreme pH environments (pH = 1) for more than 7 days, owing to its strong secondary covalent network ^[23][66]. In another study, an acid-tolerant hydrogel bioadhesive integrates two distinct components: poly(2-hydroxyethyl methacrylate-co-N-vinylpyrrolidone) (poly(HEMA-NVP)) and poly(acrylic acid-co-N-hydroxysuccinimide acrylate ester) (poly(AA-NHS)) ^[3][53]. Poly(HEMA-NVP) contributes to the acid tolerance of the hydrogel through phase segregation facilitated by hydrophobic association, intrinsic van der Waals interactions, and hydrogen bonds. Poly(AA-NHS) functions as an adhesive brush to form robust and seamless interfacial adhesion through the dry-crosslinking mechanism. The three-layered structure effectively seals 5 mm perforation in a few seconds and remains stable for more than 14 days ^[3][53].

In addition to the formulation and structure of the hydrogel, it is also critical to consider the shape of the hydrogel as it determines the route of administration and the efficacy of perforation sealing. Microparticle-based hydrogels are easily sprayable and endoscopically deliverable but may be inappropriate for larger perforation sizes due to their lowering of burst strength ^[17][28]. Glue-type hydrogels are suitable for minimally invasive surgeries via the endoscope or laparoscope. However, the hydrogel solution must show good injectability, quick gelation, and ease of operation. A multi-step gelation process that involves light administration and extra reactants may require specially designed delivery devices ^[4][24][54,67]. Patch-type hydrogels are designed for perforation sealing during open surgery and should stick instantly upon application. A major clinical problem of conventional adhesives is the undesired postoperative tissue adhesion due to indiscriminate adhesion, which causes severe consequences including chronic pain, ileus, and infertility ^[26][69]. To overcome this challenge, Janus hydrogels are being investigated with single-sided wet adhesion capabilities. A negatively charged carboxyl-containing hydrogel can be gradiently complexed with a cationic oligosaccharide via the one-sided dipping method to form a Janus hydrogel with excellent asymmetric adhesion and non-adhesion on two surfaces ^[26][69]. Alternatively, single-sided patterning with Fe³⁺ through a paper-based transfer printing method also affords Janus hydrogel patches ^[8][58]. Full blocking of the perforation defect helps the material anchor to the wound site completely and prevents adhesive failure. Liu et al. report a mushroom-cap inspired hyperboloid-shaped bioadhesive consisting of a dimethylacrylamide network crosslinked with N,N-cystaminebis(acrylamide) and a sodium alginate network crosslinked with calcium ions. The resulting hydrogel, casted into thin sheets and rolled to build a multilayer hyperboloid cap-stick-shaped device with an onion-like structure, facilitates endoscopic delivery and self-expandable full layer blocking ^[2][52].

Almost all published hydrogel designs show good biocompatibility, complete sealing of the perforation, and superior wound healing effects compared to conventional sutures or fibrin glue in in vivo gastric perforation models, ranging from mice and rats to rabbits and pigs. The application of the adhesive hydrogel alone accelerates the transition from inflammation to proliferation, suppresses excessive fibrosis to promote ECM remodeling, and provides nutrition for angiogenesis and re-epithelization ^{[3][7][9][12][13][15][24]}[39,53,57,59,60,62,67]. Additionally, small molecular drugs and growth factors are incorporated into the hydrogel for sustained release ^[2][7][52,57]. Fibroblast growth factors (FGFs) stimulate endothelial cell proliferation and angiogenesis in pathological conditions ^[27][70]. Both acidic FGF (FGF-1) and basic FGF (FGF-2), incorporated in hydrogels, promote further cell proliferation and migration, thereby accelerating abdominal wall repair ^[2][11][38,52]. Vonoprazan fumarate (VF), a potassium-competitive acid blocker, loaded in hydrogels, regulates intragastric pH to promote the healing of lesions ^[2][52]. Gastric acid interferes with platelet activation and deteriorates blood clots. VF inhibits the exchange of H⁺ and K⁺ in the gastric mucosa, leading to a higher local pH, a decreased risk of bleeding, and better mucosa healing ^[2][52].

An ideal hydrogel adhesive for GI perforation needs to exhibit high underwater adhesive strength to effectively seal the perforation, strong mechanical properties to prevent adhesive and cohesive failure, long-term extreme pH tolerance to remain stable in the GI environment, biocompatibility to ensure safety, and usability for ease of application. To ensure strong wet adhesion, researchers maximize the physical, chemical, or topological interactions between the hydrogel and the tissue surface through various strategies including: incorporation of a large numbers of physical or chemical anchoring moieties into the hydrogel network; elimination of interfacial water via the use of hydrophobic residues and dry hydrophilic scaffolds; and implementation of multiple adhesive mechanisms in parallel. To realize acid tolerance, supramolecular assemblies through hydrophobic interactions, nonhydrolyzable double covalent networks, and pH-responsive acid-resisting drugs are being explored for novel hydrogel designs. To achieve enhanced toughness, strength, and ductility, biomimetic TA-based hydrogels and IPN hydrogels with both covalent and physical crosslinking networks provide both cohesion and adhesion. Additionally, hydrogels in different forms are being investigated to accommodate various surgical delivery routes and perforation sizes. Compared to conventional sutures and bioadhesive glues, hydrogel adhesives generally show high perforation sealing efficacies and superior wound healing effects in in vivo animal models. Many formulations show promise for clinical translation. However, due to the complexity of the disease, a universal hydrogel adhesive solution for GI perforation treatment seems unlikely. Additionally, one should be cognizant of industrial scalability and the companion device needed for delivery and not create an overly complex hydrogel adhesive. Significant opportunities exist to advance better hydrogel designs and treatment outcomes for GI perforation.