WO2025015266A1

WO2025015266A1 - A highly elastic and tough hydrogel for efficient hemostasis and closure of injured internal organs in emergency healthcare

Info

Publication number: WO2025015266A1
Application number: PCT/US2024/037781
Authority: WO
Inventors: Nasim Annabi; Yuting Zheng
Original assignee: University of California Berkeley; University of California San Diego UCSD
Current assignee: University of California Berkeley; University of California San Diego UCSD
Priority date: 2023-07-12
Filing date: 2024-07-12
Publication date: 2025-01-16
Anticipated expiration: 2026-01-12

Abstract

Disclosed is a novel strategy for engineering tissue adhesives in which molecular building blocks are manipulated to allow for precise control and optimization of the various aforementioned properties without any tradeoffs. To introduce tunable mechanical properties and robust tissue adhesion, the hydrogel network presents different modes of covalent and noncovalent interactions using N-hydroxysuccinimide ester (NHS) conjugated alginate (Alg-NHS), poly (ethylene glycol) diacrylate (PEGDA), tannic acid (TA), and Fe3+ ions. Through combining and tuning different molecular interactions and a variety of crosslinking mechanisms, we were able to design an extremely elastic (924%) and tough (4697 kJ/m3) multifunctional hydrogel that could quickly adhere to wet tissue surfaces within 5 sec of gentle pressing and deform to support physiological tissue function over time under wet conditions.

Description

A HIGHLY ELASTIC AND TOUGH HYDROGEL FOR EFFICIENT HEMOSTASIS AND CLOSURE OF INJURED INTERNAL ORGANS IN EMERGENCY HEALTHCARE

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims the benefit under 35 U.S.C. Section 119(e) of copending and commonly-assigned U.S. Provisional Patent Application No. 63/513,232, filed July 12, 2023, entitled “A HIGHLY ELASTIC AND TOUGH HYDROGEL FOR EFFICIENT HEMOSTASIS AND CLOSURE OF INJURED INTERNAL ORGANS IN EMERGENCY HEALTHCARE”, the contents of which is incorporated by reference herein.

STATEMENT REGARDING FEDERAL FUNDING

This invention was made with government support under EB023052, and HL 140618 awarded by the National Institutes of Health. The government has certain rights in the invention.

TECHNICAL FIELD

Embodiments of the disclosure concern at least the fields of medicine and material science.

BACKGROUND OF THE INVENTION

Minimally-invasive sealing of injured tissues and organs is of utmost importance in biomedicine. The global market for hemostats and tissue sealants was more than $4.6 billion in 2017, which is expected to grow beyond $10 billion by 2027. The integration of sealant biomaterial with tissues undergoing extensive daily activities, such as stretching and contraction, is an unmet clinical need that requires stretchable and biocompatible material platforms. Various tissue injuries, such as skin and muscle laceration and burn, cardiac trauma, lung puncture, liver bleeding, cornea, and teeth and gum injuries as well the anastomosis of tubular tissues have been treated with tissue adhesive biomaterials to leverage a regenerative microenvironment while providing a physical support. Controlling traumatic bleeding from damaged internal organs while effectively sealing the wound is critical for saving the lives of patients. Unfortunately, conventional bioadhesive materials suffer from a number of drawbacks including blood incompatibility, insufficient adhesion to wet surfaces, weak mechanical properties, and complex application procedures. There is currently a significant medical need for new bioadhesives and tough tissue sealant materials that can, for example, control traumatic bleeding from damaged internal organs.

SUMMARY OF THE INVENTION

As disclosed herein, we have designed a ready-to-use hemostatic bioadhesive with ultra-strengthened mechanical properties and fatigue resistance, robust adhesion to wet tissues within a few seconds of gentle pressing, deformability to accommodate physiological function and action, and the ability to stop bleeding efficiently. The engineered hydrogel, which demonstrated high elasticity (> 900%) and toughness (>4600 kJ/m3), was formed by fine-tuning a series of molecular interactions and crosslinking mechanisms involving N-hydroxysuccinimide (NHS) conjugated alginate (Alg-NHS), poly (ethylene glycol) diacrylate (PEGDA), tannic acid (TA), and Fe3+ ions. Dual adhesive moieties including mussel-inspired pyrogallol/catechol and NHS synergistically enhanced wet tissue adhesion (> 400 kPa in a wound closure test). In conjunction with physical sealing, the high affinity of TA/Fe3+ for blood could further augment hemostasis. The engineered bioadhesive demonstrated excellent in vitro and in vivo biocompatibility as well as improved hemostatic efficacy as compared to the commercial material Surgicel®. Moreover, the hydrogel design strategy described herein can be used to overcome existing obstacles that are impeding the clinical translation of engineered hemostatic bioadhesives.

The invention disclosed herein has a number of embodiments, for example, methods of synthesizing the hydrogels disclosed herein as well as method for using them. In illustrative working embodiments of the invention, Alg-NHS and PEGDA were first separately synthesized as the primary molecular backbone of the resulting APTF hydrogel. NHS was conjugated to Alg to provide instant tissue-material interfacial interactions and to fine-tune the ionic interactions with Fe3+ to control the mechanical properties. TA was also introduced to enrich the crosslinking network by forming dynamic H-bonding interactions with PEGDA, and ion- induced chelation with Fe3+ to further tune the mechanical properties while providing multiple functions (i.e., antioxidant and antibacterial properties) for the hydrogel. Additionally, the mussel-inspired adhesive characteristics of TA, along with NHS functionalization onto Alg, provided enhanced tissue adhesion to the resulting APTF hydrogel. The covalent and noncovalent interactions among various macromolecules in the designed hydrogel were used to tune its physical properties. Specifically, the mixture of Alg- NHS/PEGDA prepolymer solution was treated under visible blue light (405 nm), in the presence of lithium phenyl-2,4,6- trimethylbenzoylphosphinate (LAP) photoinitiator to form a covalently interconnected PEGDA network interpenetrated with Alg-NHS, making the primary hydrogel network (AP hydrogel). A secondary molecular network was then established with the addition of TA molecules, which facilitated the formation of H-bonding between oxygen-rich PEGDA backbones. Meanwhile, an ionic crosslinking network was formed inside the hydrogel through the incorporation of Fe3+ ions, which could interact with both TA and Alg-NHS, leading to the resulting APTF hydrogel patch an “all-in-one” crosslinked network.

As shown in the examples below, the invention disclosed herein has a number of embodiments. Embodiments of the invention include the hydrogel compositions disclosed herein, typically those comprising N-hydroxysuccinimide ester (NHS) conjugated alginate (Alg-NHS); poly (ethylene glycol) diacrylate (PEGDA); tannic acid (TA); and Fe³⁺ ions. Optionally such compositions include a photoinitiator agent. In certain embodiments of the invention, the composition includes a therapeutic agent, an imaging agent or the like. Typically in such compositions, the hydrogel exhibits at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa. Embodiments of the invention also include methods of making hydrogel compositions, typically by combining an alginate conjugated to N- hydroxysuccinimide ester moieties with a poly (ethylene glycol) diacrylate (PEGDA); including in the combination tannic acid (TA); and Fe³⁺ ions; such that the hydrogel composition is made. Typically in these methods, the hydrogel is formed to exhibit at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa. Some of these methods further comprise including in the combination a photoinitiator agent, a therapeutic agent, an imaging agent or the like. Certain of these methods comprise exposing the composition to visible blue light so as to activate the photoinitiator agent and form a covalently interconnected PEGDA network interpenetrated with Alg-NHS.

Embodiments of the invention include methods of delivering a composition disclosed herein to a preselected site, the methods comprising disposing the composition at the site and then crosslinking the composition in situ. Typically in these methods, the site is an in vivo .site such as an in vivo location where an individual has experienced trauma or injury. In illustrative embodiments of these methods, the site comprises vascular, bladder, lung or heart tissue.

Related embodiments of the invention include methods of adhering a first patch material and/or a first tissue region to a second tissue region. Typically these methods comprise disposing a hydrogel composition disclosed herein at a site where the hydrogel composition is in contact with the first tissue region and the second tissue region; and forming a covalently interconnected polymer network form the hydrogel composition; such that the first tissue region is adhered to the second tissue region. In some embodiments of the invention, the first tissue region is adhered to the second tissue region in vivo. Certain of these methods comprise exposing the composition to visible blue light so as to activate the photoinitiator agent and form a covalently interconnected network. In some methods of the invention, the hydrogel composition is prefabricated as an adhesive patch. In some methods of the invention, the hydrogel composition is 3D printed. In some methods of the invention, the hydrogel composition inhibits bleeding as the first tissue region is adhered to the second tissue region. In some methods of the invention, the hydrogel composition further comprises a mammalian cell, a therapeutic agent or an imaging agent.

The methods and materials disclosed herein have a number of advantages over existing bioadhesive sealants. By tuning different molecular interactions in methods of making bioadhesive embodiments of the invention, a single-hydrogel platform with an “all-in-one” multifunctionality can be engineered. Moreover, the bioadhesive embodiments of the invention possess selectively tunable mechanical properties matching the stiffness of different types of tissue. Ultra- strengthened toughness, superior elasticity, high strength, and fatigue-resistance capability can be achieved. Bioadhesive embodiments of the invention can maintain mechanical integrity over time under wet conditions and demonstrate minimal swelling. Such bioadhesive embodiments of the invention further display antioxidant and antibacterial properties necessary for facilitating the wound healing process, and are biodegradable. The bioadhesive embodiments of the invention also demonstrated robust wet adhesion (strong adhesion forms in 5s of gentle pressing) and painless detachable characteristics, meanwhile ensuring excellent in vitro and in vivo biocompatibility.

Bioadhesive embodiments of the invention can be prefabricated as an adhesive patch and ready-to-use in emergency healthcare, extremely easy to apply and can store up to 6 months prior to usage without losing its adhesion. The bioadhesive embodiments of the invention can achieve long-term adhesion without detachment. In this way, bioadhesive embodiments of the invention can rapidly stop bleeding while effectively sealing injured tissue. In addition, bioadhesive embodiments of the invention can be 3D printed into different sophisticated structures for use in more complex and challenging biomedical applications. Detailed characterizations and data illustrating these features can be found in the Examples below.

Other objects, features and advantages of the present invention will become apparent to those skilled in the art from the following detailed description. It is to be understood, however, that the detailed description and specific examples, while indicating some embodiments of the present invention, are given by way of illustration and not limitation. Many changes and modifications within the scope of the present invention may be made without departing from the spirit thereof, and the invention includes all such modifications.

BRIEF DESCRIPTION OF THE DRAWINGS

Fig. 1. Characterizations of synthesized Alg-NHS, AP hydrogel, APTF hydrogel, (a) Schematic for Alg-NHS synthesis, (b) 'H NMR and (c) FTIR spectra of Alg, Alg conjugated with NHS under different reaction conditions. FTIR spectra showing (i) carbonyl stretch of NHS (1780 cm'¹), (ii) carbonyl stretch of NHS (1704 cm'¹), (iii) CNC stretch of NHS (1219 cm'¹), (d) The viscosity of synthesized Alg- NHS and Alg at 2%. (e) Rheology characterizations (G’, G”) of Alg/Fe³⁺, Alg- NHS/Fe³⁺, and Alg-NHS/TA/Fe³⁺ hydrogel, the concentrations of Alg, Fe³⁺ and TA in this study were 2%, 3% and 40% respectively, (f) FTIR spectra of photocrosslinked PEGDA hydrogel (22%) and AP hydrogel consisting of 8% Alg-NHS and 22% PEGDA: (i) carbonyl stretch of NHS (1780 cm'¹), (ii) carbonyl stretch of NHS (1704 cm'¹), (iii) CNC stretch of NHS (1219 cm'¹), (g) XPS survey of PEDGA and APTF hydrogel, (h) Carbon XPS of APTF hydrogel.

Fig. 2. Physical characterizations of PEGDA, PEGDA/TA, AP, PTF, Alg- NHS/PEGDA hydrogel with treated with TA (APT) and APTF hydrogel, (a) Macroscopic images of PEGDA hydrogels crosslinked with TA (40%), and TA/Fe³⁺ with varied Fe³⁺ concentrations (3%, 6%). (b) Releasing profile of TA from PEGDA/TA or PTF hydrogels with varied Fe³⁺ concentrations, (c) UV-Vis of released TA from PEGDA/TA or PTF hydrogels, (d) SEM images of the (i) PEGDA, (ii) AP, (iii) PEGDA/TA, (iv) PTF and (v) APTF hydrogel. Scar bar = 10 pm.

Fig. 3. Effects of treatment methods on mechanical properties of APTF hydrogel, (a) Young’s modulus, (b) strain, (c) ultimate strength, and (d) toughness of APTF hydrogels using different treatment methods for crosslinking. We characterized the tunability of the hydrogel’s mechanical properties, which was predominantly controlled by creation or alteration of the molecular interactions and crosslinking nature across the network. In our system, Alg-NHS interacts with Fe³⁺ through ionic interactions while Fe³⁺ chelates with TA in a pH-sensitive manner. Fe³⁺ forms bis-chelation with TA when the pH is higher than 1.7 (88). In this work, the observed chelation was a mono-chelation as the pH of the 40% TA solution was highly acidic (pH =1.5). This was further confirmed by ultraviolet-visible spectroscopy (UV-vis) which exhibited no characteristic bis-chelation and trischelation peaks around 600 nm (89).

Fig. 4. Mechanical characterizations of APTF hydrogels formed by varying Alg-NHS/PEGDA ratio, (a) Mechanical properties of APTF hydrogels with fixed total polymer concentration (22%) and varied treatment methods (40% TA/0%, 3%, 6% Fe³⁺). (i) Young’s modulus, (ii) ultimate strength and (iii) toughness and (iv) ultimate strain of hydrogels treated with TA/ Fe³⁺. Young’s modulus, strain, and ultimate strength of APTF hydrogels formed with varied (b) Alg-NHS concentrations and (c) PEGDA concentrations, (d) Toughness of APTF hydrogels with varied Alg- NHS concentrations, (e) Representative stress-strain curve of the optimized engineered APTF hydrogel. (1) Stretching of the APTF hydrogel, (g) Mechanical parameters of optimized APTF hydrogel.

Fig. 5. Adhesion assessment of APTF hydrogels, (a) Illustration of adhesion mechanism to wet tissue surfaces, (b) Shear strength of different hydrogels on porcine skin and their (c) typical lap shear stress-strain profiles, (d) Demonstration of adhesiveness of APTF patch to different tissues (one side adhered to a spatula and another side adhered to tissue within 5 s gentle pressing), (e) Adhesive strength of APTF hydrogel using two ASTM testing methods including lap shear test and tensile pull-off test. (1) Shear strength of APTF hydrogel on different wet tissues, (g) Representative lap shear stress-strain curve and (h) shear strength of APTF hydrogels on porcine skin with and without urea or DFO treatment, (i) Shear strength of APTF hydrogels on porcine cornea and conjunctiva with and without urea treatment, (j) Demonstration of strong adhesion to rabbit conjunctiva in situ.

Fig. 6. In vitro biocompatibility and antioxidant activity of APTF hydrogel. In vitro cell studies: (a) Representative live/dead stained images of 3T3 cells on AP hydrogel and APTF hydrogel on day 1 and day 7. (b) Quantification of cellular viability for 3T3 cells seeded on AP and APTF hydrogel over 7 days of culture, (c) Cellular proliferation on AP hydrogel and APTF hydrogel over time based on PrestoBlue assay. In vitro antioxidant activity: (d) The color changes of the three DPPH* solutions containing hydrogel-lacking control, AP hydrogel, APTF hydrogel over time, (e) Absorbance change of DPPH* before and after the reaction, (f) The DPPH* scavenging activities of AP and APTF hydrogels.

Fig. 7. Photothermal and antibacterial properties of APTF hydrogels. Temperature enhancement over time for the (a) dry APTF hydrogels and (b) wet APTF hydrogels under irradiation with a NIR laser (808 nm) under different power densities, (c) Photothermal stability of the engineered APTF hydrogels undergoing 10 consecutive heating cycles, (d) Infrared thermal images of the AP and APTF hydrogels under 808 nm irradiation for 6 min. (e) Quantitative antibacterial efficiency against p. aeruginosa under different treatment conditions. (1) Quantitative antibacterial efficiency of APTF hydrogel against p. aeruginosa and MRSA under 808 nm irradiation, (g) Representative images from the p. aeruginosa colonies on the agar plates after treatment with DPBS, DPBS+light, AP hydrogel +light, APTF hydrogel, APTF hydrogel+light. Images were taken from a Zeiss Axio Observer Z1 inverted microscope. Scar bar = 500 pm. (h) Cell viability of APTF hydrogel under 808 nm irradiation.

Fig. 8 Scheme 1. (a) Schematically illustrated synthesis of APTF hydrogel based on Alg-NHS, PEGDA, and TA/Fe³⁺. (b) Molecular interactions among the building blocks of APTF hydrogel: (i) TA crosslinked PEGDA network (ii) TA/Fe³⁺ crosslinked PEGDA network and (iii) TA/Fe³⁺ crosslinked AP network. FIG. 9. Scheme 1. Another view of the schematically illustrated synthesis of APTF hydrogel based on Alg-NHS, PEGDA, and TA/Fe³⁺ as shown in FIG. 8.

FIG 10. Mechanical and physical characterizations of the APTF hydrogel, (a) Schematics of molecular interactions (i-iv) and the corresponding SEM images (v- viii) among the building blocks of the APTF hydrogel: (i) and (v) denote the PEGDA network, (ii) and (vi) indicate the TA crosslinked PEGDA network, (iii) and (vii) show the TA/Fe³⁺ crosslinked PEGDA network, (iv) and (viii) stand for TA/Fe³⁺ crosslinked AP network, (b-c): Young’s modulus and ultimate strength of APTF hydrogels (40% TA/3% Fe³⁺ treatment) formed with varied (b) Alg-NHS concentrations and (c) PEGDA concentrations, (d) Photographs demonstrating the strong mechanical properties and elasticity of the APTF hydrogel, (e-f): Strain and toughness of APTF hydrogels (40% TA/3% Fe³⁺ treatment) formed with varied (e) Alg-NHS concentrations and (1) PEGDA concentrations. Error bars indicate standard error of the means, asterisks mark significance levels of p < 0.05 (*), p < 0.001 (***), p < 0.0001 (****) and n >5.

FIG 11. Mechanical and physical characterizations of the APTF hydrogel, (a) Sequential loading-unloading compression tests without interval and (b) the corresponding calculated total and dissipated toughness of the APTF hydrogel under different strains (e = 10%, 20%, 50% and 60%). (c) Hysteresis loops of cycle 1, 50,100 of the APTF hydrogel, (d) Compressive toughness and dissipated toughness of the APTF hydrogel over 100 cyclic compression tests, (e) Literature review comparing mechanical properties of TA-based adhesive hydrogels versus the APTF hydrogel. The pink region denotes hydrogels with high brittleness featuring low mechanical strength and stretchability, the orange region contains those with high stretchability but low stiffness and ultimate strength. The region in blue shows a midpoint between the two previously mentioned regions comprised of middling stretchability and mediocre toughness, (f) Swelling profile and representative images of the APTF hydrogel and AP hydrogel, (g) Photograph of customizable shapes and structures of APTF hydrogels formed through PDMS molds, (h) Demonstration of layer-by-layer printing of AP bioink loaded with food dye (i.e., green and blue), (i) Demonstration of mechanical properties of 3D printed AP hydrogel and APTF hydrogel post TA/Fe³⁺ treatment, (j) The viscosity of synthesized AP bioink at different Alg-NHS concentrations (8% and 12%) as a function of shear rate, (k) Demonstration of the 3D printed tubular structure: (i) printing of AP bioink and (ii) crosslinking the printed structure with UV light (402 nm) to form (iii) a one cm thickness layer-by-layer printed tubular construct. Error bars indicate standard error of the means, asterisks mark significance levels of p < 0.05 (*), p < 0.001 (***), p < 0.0001 (****) and n =4 for the cyclic compression test, n =5 for the swelling test.

FIG 12. In vitro and ex vivo adhesive characterizations of the APTF hydrogel, (a) Shear strength of APTF hydrogels with different compositions (40% TA treatment with and without adding 3% Fe³⁺). (b) Shear strength of APTF hydrogels with increased PEGDA concentrations (40% TA/3% Fe³⁺ treatment) in comparison with the commercial cyanoacrylate glue, (c) Illustration of strong adhesiveness of the APTF hydrogel to the porcine skin and (d) its adhesiveness underwater, (e) Adhesive strength and interfacial toughness of the APTF hydrogel using three ASTM testing methods (lap shear test, 180-degree peel test and wound closure test) and their (f) illustrations, (g) Shear strength of the APTF hydrogel on different tissues, (h) The effect of hydrogel incubation time on the shear strength of the APTF hydrogel, (i) Shear strength of the APTF hydrogel on different substrates, (j) Demonstration of the ex vivo burst pressure tests on (i) heart (5 mm laceration), (ii) intestine (1.5 mm-diameter circular incision), (iii) aorta (5 mm laceration) and ex vivo sealing demonstration of the APTF hydrogel on (iv) bladder (1.5 mm circular incision) and (v) stomach (10 mm-diameter circular incision), (k) Burst pressure values of different organs sealed with APTF hydrogels. (1) Schematic demonstrating non-covalent and covalent interactions between the tissue surfaces and the APTF hydrogel. Error bars indicate standard error of the means, asterisks mark significance levels of p < 0.05 (*), p < 0.001 (***), p < 0.0001 (****) and n > 5. FIG 13. In vitro and in vivo biocompatibility of the APTF hydrogel, (a) Representative live/dead images of 3T3 cells seeded on the underside of the Transwell permeable supports with APTF hydrogels and cyanoacrylate (CA) 1- and 5-days postseeding (scale bars = 100 pm), (b) Quantification of cellular viability over 5 days of culture, (c) Representative F-actin/DAPI stained images of 3T3 cells seeded the underside of the Transwell permeable supports with APTF hydrogels 1- and 5-days post-seeding (scale bars = 50 pm), (d-e) Quantitative analysis of cellular metabolic activity, relative fluorescence units (RFU) of (d) cyanoacrylate and (e) APTF hydrogel using a PrestoBlue assay at days 1 and 5 post-seeding, (f) Hematoxylin and eosin (H&E), (g) Masson’s tri chrome (MT), and (h) immunofluorescence (IF) staining from the APTF hydrogel/tissue interfaces at days 7 and 28 post-implantation (scale bars = 100pm). Anti-CD68 (abl25212) (Abeam) was utilized as a primary antibody, while Goat anti-Rabbit IgG (H+L) antibody conjugated to Alexa Fluor 488 (Invitrogen) was used as a detection reagent and secondary antibody. Cell infiltration inside the implanted hydrogels is marked with yellow arrows. The subcutaneous implantation model in Wistar rats was chosen to assess the biocompatibility of the engineered APTF patch in vivo. The types of cells found in subcutaneous tissue are fibroblasts, adipose cells, and macrophages. Error bars indicate standard error of the means, asterisks mark significance levels of p < 0.05 (*), p < 0.001 (***), p < 0.0001 (****) and n = 4.

FIG 14. In vitro and in vivo hemostatic properties of the APTF hydrogel, (a) Representative images of clot formation after exposing the hydrogel formations to the fresh blood and (b) the in vitro blood clotting formation kinetic profile to evaluate the hemostasis of AP and APTF hydrogels, (c-e) In vivo rat tail bleeding model to study the hemostatic efficacy of the hydrogel: (c) Images of the rat tail bleeding surgical procedure, (d) the mass of collected blood on the filter paper following treatment with AP hydrogel, APTF hydrogel, and Surgicel® as compared with the injury group with no treatment and (e) representative images of the blood collected on the filter papers for different treatment groups in the rat tail bleeding model, (f-g) In vivo liver bleeding model to study the hemostatic efficacy of the hydrogel: (f) images of the rat liver bleeding surgical procedure, and (g) the mass of collected blood on the filter paper following treatment with AP hydrogel, APTF hydrogel, and Surgicel® as compared with the injury group with no treatment. Error bars indicate standard error of the means, asterisks mark significance levels of p < 0.05 (*), p < 0.001 (***), p < 0.0001 (****) and n > 4.

DETAILED DESCRIPTION OF THE INVENTION

In the description of embodiments, reference may be made to the accompanying figures which form a part hereof, and in which is shown by way of illustration a specific embodiment in which the invention may be practiced. It is to be understood that other embodiments may be utilized, and structural changes may be made without departing from the scope of the present invention. Many of the techniques and procedures described or referenced herein are well understood and commonly employed by those skilled in the art. Unless otherwise defined, all terms of art, notations and other scientific terms or terminology used herein are intended to have the meanings commonly understood by those of skill in the art to which this invention pertains. In some cases, terms with commonly understood meanings are defined herein for clarity and/or for ready reference, and the inclusion of such definitions herein should not necessarily be construed to represent a substantial difference over what is generally understood in the art.

Designing adhesive hydrogels with optimal properties for the treatment of injured tissues is challenging due to the tradeoff between material stiffness and toughness, while maintaining adherence to wet tissue surfaces. In most cases, bioadhesives with improved mechanical strength, often lack an appropriate elastic compliance, hindering their application for sealing soft, elastic, and dynamic tissues. Here, we present a novel strategy for engineering tissue adhesives in which molecular building blocks are manipulated to allow for precise control and optimization of the various aforementioned properties without any tradeoffs. To introduce tunable mechanical properties and robust tissue adhesion, the hydrogel network presents different modes of covalent and noncovalent interactions using N- hydroxysuccinimide ester (NHS) conjugated alginate (Alg-NHS), poly (ethylene glycol) diacrylate (PEGDA), tannic acid (TA), and Fe³⁺ ions. Through combining and tuning different molecular interactions and a variety of crosslinking mechanisms, we were able to design an extremely elastic (924%) and tough (4697 kJ/m³) multifunctional hydrogel that could quickly adhere to wet tissue surfaces within 5 sec of gentle pressing and deform to support physiological tissue function over time under wet conditions. While Alg-NHS provides covalent bonding with the tissue surfaces, the catechol moieties of TA molecules synergistically adopt a mussel-inspired adhesive mechanism to establish robust adherence to the wet tissue. The strong adhesion of the engineered bioadhesive patch is showcased by its application to rabbit conjunctiva and porcine cornea. Meanwhile, the engineered bioadhesive demonstrated painless detachable characteristics and in vitro biocompatibility. Additionally, due to the molecular interactions between TA and Fe³⁺, antioxidant and antibacterial properties required to support the wound healing pathways were also highlighted. Overall, by tuning various molecular interactions, we were able to develop a singlehydrogel platform with an “all-in-one” multifunctionality that can address current challenges of engineering hydrogel-based bioadhesives for tissue repair and sealing.

In addition, controlling traumatic bleeding from damaged internal organs while effectively sealing the wound is critical for saving the lives of patients. Existing bioadhesives suffer from blood incompatibility, insufficient adhesion to wet surfaces, weak mechanical properties, and complex application procedures. Here, we engineered a ready -to-use hemostatic bioadhesive with ultra- strengthened mechanical properties and fatigue resistance, robust adhesion to wet tissues within a few seconds of gentle pressing, deformability to accommodate physiological function and action, and the ability to stop bleeding efficiently. The engineered hydrogel, which demonstrated high elasticity (> 900%) and toughness (>4600 kJ/m³), was formed by fine-tuning a series of molecular interactions and crosslinking mechanisms involving N-hydroxysuccinimide (NHS) conjugated alginate (Alg-NHS), poly (ethylene glycol) diacrylate (PEGDA), tannic acid (TA), and Fe³⁺ ions. Dual adhesive moi eties including mussel-inspired pyrogallol/catechol and NHS synergistically enhanced wet tissue adhesion (> 400 kPa in a wound closure test). In conjunction with physical sealing, the high affinity of TA/Fe³⁺ for blood could further augment hemostasis. The engineered bioadhesive demonstrated excellent in vitro and in vivo biocompatibility as well as improved hemostatic efficacy as compared to commercial Surgicel®. Overall, the hydrogel design strategy described herein holds great promise for overcoming existing obstacles impeding clinical translation of engineered hemostatic bioadhesives.

As shown in the examples below, the invention disclosed herein has a number of embodiments. Embodiments of the invention include hydrogel compositions, typically those comprising N-hydroxy succinimide ester (NHS) conjugated alginate (Alg-NHS); poly (ethylene glycol) diacrylate (PEGDA); tannic acid (TA); and Fe³⁺ ions. Typically such compositions include a photoinitiator agent. In certain embodiments of the invention, the composition includes a therapeutic agent, an imaging agent or the like. Typically in such compositions, the hydrogel exhibits at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa.

Embodiments of the invention also include methods of making hydrogel compositions, typically by combining an alginate conjugated to N- hydroxysuccinimide ester moieties with a poly (ethylene glycol) diacrylate (PEGDA); including in the combination tannic acid (TA); and Fe³⁺ ions; such that the hydrogel composition is made. Some of these methods further comprise including in the combination a photoinitiator agent, a therapeutic agent or an imaging agent or the like. Certain of these methods comprise exposing the composition to visible blue light so as to activate the photoinitiator agent and form a covalently interconnected PEGDA network interpenetrated with Alg-NHS. In some of these methods, the hydrogel is formed to exhibit at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa.

As scaffolds, compositions of the invention can provide a flexible dwelling space for cells and other agents for use in tissue repair and the regeneration of desired tissues (e.g. for cartilage, bone, retina, brain, and, neural tissue repair, vascular regeneration, wound healing and the like). Moreover, embodiments of the invention can include immunomodulatory agents useful for immunotherapy in order to, for example, enhance components of the immune system. Certain illustrative materials and methods that can be adapted for use in such embodiments of the invention are found, for example in Hydrogels: Design, Synthesis and Application in Drug Delivery and Regenerative Medicine 1st Edition, Singh, Laverty and Donnelly Eds; and Hydrogels in Biology and Medicine (Polymer Science and Technology) UK ed. Edition by J. Mi chai ek et al.

In certain embodiments of the invention, the amounts or ratios of one or more constituents is controlled in order to, for example, control one or more material properties of the compositions. In embodiments of the invention that include methods of making a composition disclosed herein, for example by combining together N- hydroxysuccinimide ester (NHS) conjugated alginate (Alg-NHS); poly (ethylene glycol) diacrylate (PEGDA); tannic acid (TA); and Fe³⁺ ions, a pharmaceutical excipient and/or a therapeutic agent so as to form the composition. By modulating the mechanical properties of the compositions of the invention in this manner, embodiments of the invention can be tailored for use in a variety of different clinical applications. In this context, a wide variety of art accepted materials and methods can be adapted for use in embodiments of the invention, for example those disclosed in U.S. Patent Publication Nos.: 20050227910, 20100120149, 20120315265, 20140302051 and 20190290804; and Lee, Biomaterials Research volume 22, Article number: 27 (2018); Thambi et al., J Control Release. 2017 Dec 10;267:57-66. doi: 10.1016/j.jconrel.2017.08.006. Epub 2017 Aug 4; and Gianonni et al., Biomater. Sci., 2016, the contents of which are incorporated by reference. Embodiments of the invention also include adhering a first tissue region to a second tissue region. Typically these methods comprise disposing a hydrogel composition disclosed herein at a site where the hydrogel composition is in contact with the first tissue region and the second tissue region; and forming a covalently interconnected polymer network form the hydrogel composition; such that the first tissue region is adhered to the second tissue region. In some embodiments of the invention, the first tissue region is adhered to the second tissue region in vivo. In some methods of the invention, the hydrogel composition is prefabricated as an adhesive patch. In some methods of the invention, the hydrogel composition is 3D printed. In some methods of the invention, the hydrogel composition inhibits bleeding as the first tissue region is adhered to the second tissue region. In some methods of the invention, the hydrogel composition further comprises a mammalian cell, a therapeutic agent or an imaging agent.

Related embodiments of the invention include methods of delivering a composition disclosed herein to a preselected site, the methods comprising disposing the composition at the site and then crosslinking the composition in situ. Typically in these methods, the site is an in vivo .site such as an in vivo location where an individual has experienced trauma or injury. In illustrative embodiments of these methods, the site comprises vascular, bladder, lung or heart tissue.

In certain embodiments of the invention, the compositions of the invention include additional constituents. In certain embodiments the compositions of the invention include one or more therapeutic agents such as an anti-inflammatory agent, an agent that modulates coagulation, an antibiotic agent, a chemotherapeutic agent or the like. Compositions of the invention can be formulated for use as carriers or scaffolds of therapeutic agents such as drugs, cells, proteins, and other bioactive molecules. Certain embodiments of the compositions of the invention include, for example a pharmaceutical excipient such as one selected from the group consisting of a preservative, a tonicity adjusting agent, a detergent, a viscosity adjusting agent, a sugar and a pH adjusting agent. For compositions suitable for administration to humans, the term "excipient" is meant to include, but is not limited to, those ingredients described in Remington: The Science and Practice of Pharmacy, Lippincott Williams & Wilkins, 21st ed. (2006) the contents of which are incorporated by reference herein.

As carriers, compositions of the invention can incorporate the agents and deliver them to a desired site in the body for the treatments of a variety of pathological conditions. In addition, as scaffolds, compositions of the invention can provide a flexible dwelling space for cells and other agents for use in tissue repair and the regeneration of desired tissues (e.g. for cartilage, bone, retina, brain, and, neural tissue repair, vascular regeneration, wound healing and the like). Moreover, embodiments of the invention can include immunomodulatory agents useful for immunotherapy in order to, for example, enhance components of the immune system. Certain illustrative materials and methods that can be adapted for use in such embodiments of the invention are found, for example in Hydrogels: Design, Synthesis and Application in Drug Delivery and Regenerative Medicine 1st Edition, Singh, Laverty and Donnelly Eds; and Hydrogels in Biology and Medicine (Polymer Science and Technology) UK ed. Edition by J. Michalek et al.

EXAMPLES

EXAMPLE 1: MOLECULAR DESIGN OF AN ULTRA-STRONG TISSUE ADHESIVE HYDROGEL WITH TUNABLE MULTIFUNCTIONALITY

Aspects of the following disclosure in Example 1 were also published in the journal Bioactive Materials, Zheng et al., Bioactive Materials: Volume 29, November 2023, Pages 214-229. (hereinafter “Zheng Bioactive Materials”).

Hydrogels, water-swollen polymeric crosslinked networks, are capable of interacting with the tissues through physical or chemical interactions and are extensively employed in the design of bioadhesive biomaterials (1, 2). However, most bioadhesives either suffer from cohesive failure due to a lack of mechanical strength, or adhesive failure resulting from insufficient tissue-biomaterial interfacial interactions (3-7). Furthermore, incorporation of additional functionalities in these adhesive hydrogels such as antibacterial and antioxidant properties that are critical for wound healing can be challenging. Mechanistically, molecular interactions enable crosslinking among polymeric backbones, forming the primary skeleton of hydrogel constructs. These crosslinked molecular domains eventually modulate the hydrogel’s mechanical properties, including toughness, stretchability, stiffness, and others. Meanwhile, the functional groups present in these crosslinked networks introduce other important features, such as adhesive, antibacterial, and antioxidant properties (8). Therefore, the ability to tune the crosslinking density while maintaining the desired physicochemical properties is imperative in designing bioadhesive hydrogels.

Engineering bioadhesive hydrogels with optimal cohesive properties that mimic the stiffness and elasticity of native tissue’s is challenging since most of the engineered hydrogel-based bioadhesives are fragile and weak (1, 7, 9-13). In addition, there is often a tradeoff among different mechanical properties of hydrogels, compromising their applicability as a potential structural biomaterial for medical use. For example, increasing the ultimate strength often reduces the stretchability, and improving the toughness may drop the stiffness of the hydrogel-based biomaterials (7, 12, 14-20). Furthermore, most bioadhesive hydrogels reported thus far have been prone to fracture and breakage when subjected to sub-MPa tensile stress or high- extent deformation, and have displayed breaking energies around merely one percent of those of biological tissues such as cartilage (21). To address these limitations, different strategies have been implemented to modify the crosslinking mechanism of hydrogel networks, including the introduction of double networks (DN) (22), nanocomposites (5), sliding-ring hydrogels (23), resilin-based hydrogels (24), topological hydrogels (25), macromolecular crosslinked networks (26), and interpenetrating network hydrogels (27). However, achieving a high tensile strength and Young’s modulus without compromising elasticity to better reflect the nature of soft elastic tissue (i.e., lung and heart) remains an unsolved issue. For instance, Daniela et al. (28) prepared photocrosslinkable gellan gum (MeGG) hydrogels combined with an ionic crosslinking system with an ultimate strength of -800 kPa, but due to the tradeoff between ionic crosslinking and the degree of methacrylation, no improvement was achieved in Young’s modulus (< 20kPa) and strain (<80%). While the stiffness could be improved by increasing the degree of methacrylation, the strain and ultimate strength were shown to significantly decreased. Recently, a hydrogen-bond-mediated, pseudo-slide-ring networked hydrogel synthesized with carboxymethylated polyrotaxanes and polyacrylamides (PAAm) showed an enhanced Young’s modulus and ultimate strength upon increasing the number of hydrogenbonding pairs. However, the stretchability did not increase in parallel with the Young's modulus and strength, as their the process generated a rigid hydrogen-bonded network that in turn diminished the mobility of the carboxymethylated a- cyclodextrins within the crosslinked hydrogel (29). In addition, not only does the use of acrylamide moieties potentially reduce the biocompatibility (30), but also the multi-step synthesis procedure increases the complexity of the system, hindering the clinical translation of this hydrogel for surgical applications. Meanwhile, as native soft tissues exist across a wide range of mechanical properties with the Young’s modulus ranging from 3 kPa for the kidney to 100 kPa for skin (31), it is essential to achieve control over the bioadhesive’s crosslinking mechanisms and the resulting mechanical properties in order to achieve tissue-mimicking mechano-physical properties, and eliminate inflammation induced by mechanical mismatch (32). Another challenge includes maintaining the mechanical integrity of the bioadhesives under wet physiological conditions over time. The stiffness of bioadhesive hydrogels can drop due to either its unstable crosslinking resulting in hydrogel network dissociation or degradation (33, 34) or the excessive swelling of the hydrophilic network (35-39). Therefore, there is an unmet need for a biocompatible hydrogel platform featuring a programmable crosslinking mechanism which allows for fine tuning of a single material property (e.g., stiffness, toughness, stretchability, ultimate strength) without compromising another. In addition to the improved cohesion achieved by modulating the crosslinking mechanism, the wet tissue adhesive properties displayed by these biomaterials are also important for their applications as sealants or glues for sealing injured internal organs.

Specially for the closure of internal wounds, adhesive hydrogels are an alternative to sutures and staples (40-42). The retention of these bioadhesives on injured tissue surfaces throughout the healing process while supporting physiological tissue function is critical. Two main design strategies have been used to impart wet tissue adhesion to hydrogels: 1) in situ formation of adhesive hydrogels on wet tissue from precursor solutions, and 2) adhesion of prefabricated patches to the wet tissue surfaces (1, 40, 43, 44). However, both design strategies face limitations. For example, many in .w/zz-forming adhesive hydrogels require the use of an external device, such as a light source (4, 45-47) or co-injector delivery system (48, 49), while others require adding/spraying crosslinkers (50, 51) or mixing different components (52), not only further complicating the design but also making the gelation inhomogeneous. These extra steps can be fatal in urgent, life-threatening situations. In addition, traditional bioadhesives only demonstrate strong adhesion to dry substrates in air and cannot effectively seal wet organs (11, 53, 54). Slow gelation presents another hindrance for the clinical application of these hydrogels. For example, Juan et al. successfully prepared a bioadhesive prepolymer solution for in situ gelation based on tannic acid (TA) and silk fibroin, but the gelation process did not occur until after 9 h (55). In another study, Younseon et al. made a phenol-amine superglue with a high adhesive strength of 3.87 MPa, but the required curing time was 12 h (56). A slow gelation time can render the bioadhesive ineffective as it cannot prevent excessive blood loss and seal leakages or undergo application to dynamic organs such as the heart or lungs. In addition, body fluids may foul or dilute the prepolymer solution before effective gelation in situ. On the other hand, a rapid gelation process also presents complications with handling of the bioadhesive material by clinicians. In many existing studies on prefabricated adhesive patches, there is an unspecified or a prolonged incubation or contact time (>2h) necessary for the patch to achieve robust adhesion. (1, 40, 44, 57-60). Additionally, previously reported hydrogel-based adhesives demonstrate insufficient sealing strength (with most having an adhesive strength of less than 40 kPa to skin tissue) and short-term adhesion (7, 61, 62). Furthermore, current bioadhesives have shown limited characteristics in different surgical applications. For example, Xin et al prepared an ultra-fast self-gelling powder that could crosslink in situ within two sec; however, the hydrogel crosslinking was highly pH-dependent, limiting its usage for internal organs such as stomach with acidic environment (63). Moreover, it remains an obstacle to both achieve strong cohesion and adhesion simultaneously (3-7, 64), and also allow for painless detachment in situations involving misplacement, the need to adjust the material, or implanted device retrieval. To the best of our knowledge, few bioadhesives developed so far have demonstrated robust and rapid adhesion along with trigger-induced detachability (65). Therefore, there is an unmet clinical need for bioadhesives which are designed via their molecular domains to incorporate multi-mode crosslinking sites that appropriately modulate network mechanics and functionalities while also facilitating tissue-material interfacial interactions.

Incorporating multiple functionalities (e.g., antimicrobial properties) into bioadhesive hydrogels is an important step to make them suitable for both the sealing and repairing of internal wounds. For example, researchers have incorporated different antimicrobial peptides (11), antibacterial nanomaterials such as silver nanoparticles (66), and antibiotics (67) into bioadhesives to provide them with antibacterial properties. In addition, different chemical functional groups such as quaternary ammoniums (68) have been conjugated to the bioadhesives to make them antimicrobial. However, these methods require additional steps during synthesis that have often involved expensive and sensitive materials that can cause issues of cytotoxicity or bacterial resistance, and provide only short-term antibacterial effects due to leaching of antibiotics or nanomaterials out of the polymeric network (69).

To address the need for designing multifunctional adhesive biomaterials with antimicrobial properties for the treatment of different injured tissues, we aimed to develop an elastic and tough hydrogel patch that adheres to wet tissues after a few seconds of pressing while supporting the tissue’s physiological functions. This multifunctional patch, entitled APTF hydrogel, was composed of a poly (ethylene glycol) diacrylate (PEGDA) hydrogel interpenetrated with A-hydroxysuccinimide ester (NHS)-conjugated alginate (Alg-NHS), that was subsequently treated with TA/Fe³⁺ to achieve robust cohesion and adhesion. We established tunability of the mechanical and adhesive properties of the hydrogel by controlling a series of covalent and non-covalent interactions among Alg-NHS, PEGDA, TA and Fe³⁺ that collectively formed a multi-component macromolecular system. While the covalent network of PEGDA defined the primary backbone of the hydrogel, the reversible hydrogen bonding between PEGDA and TA, in addition to ionic interactions between Alg-NHS and Fe³⁺, together provided excellent elasticity and toughness. As such, stiffness, ultimate strength, and toughness could be improved significantly without compromising stretchability. The robust adhesive nature of the hydrogel was achieved by chemically conjugating NHS to alginate strands and incorporating the mussel- inspired adhesive moieties presented by TA; the synergy effect of these two adhesive moieties on adhesion enhancement was established. Strong adherence of the designed bioadhesive was demonstrated on rabbit conjunctiva and porcine cornea. Further, painless detachment of the patch was also easily achieved upon the addition of nontoxic chemical agents to the tissue-material interface. Meanwhile, the engineered hydrogel showed excellent antioxidant and antibacterial properties due to the presence of the TA/Fe³⁺ complexes that are essential for wound treatment. The biocompatibility of the multifunctional bioadhesive was assessed in vitro using 3T3 cells. Overall, through molecular engineering, the designed hydrogels demonstrated selectively tunable mechanical and physiochemical properties not only to facilitate the closure and treatment of internal and external injuries, but also to potentially promote wound healing. Synthesis of APTF adhesive hydrogel based on Alginate (Alg)-NHS, PEGDA, and TA/Fe³⁺

Sequential synthesis of the APTF adhesive hydrogel patch comprised of Alginate (Alg)-NHS, PEGDA and TA/Fe³⁺ is demonstrated in Scheme la shown in Fig. 8. First, Alg-NHS and PEGDA were separately synthesized as the primary molecular backbone of the resulting APTF hydrogel. NHS was conjugated to Alg to provide instant tissue-material interfacial interactions and to fine-tune the ionic interactions with Fe³⁺ to control the mechanical properties. TA was also introduced to enrich the crosslinking network by forming dynamic H-bonding interactions with PEGDA, and ion-induced chelation with Fe³⁺ to further tune the mechanical properties while providing multiple functions (i.e., antioxidant and antibacterial properties) for the hydrogel. Additionally, the mussel-inspired adhesive characteristics of TA, along with NHS functionalization onto Alg, provided enhanced tissue adhesion to the resulting APTF hydrogel. The covalent and noncovalent interactions among various macromolecules in the designed hydrogel were used to tune its physical properties. Specifically, the mixture of Alg-NHS/PEGDA prepolymer solution was treated under visible blue light (405 nm), in the presence of lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) photoinitiator to form a covalently interconnected PEGDA network interpenetrated with Alg-NHS, making the primary hydrogel network (AP hydrogel). A secondary molecular network was then established with the addition of TA molecules, which facilitated the formation of H- bonding between oxygen-rich PEGDA backbones. Meanwhile, an ionic crosslinking network was formed inside the hydrogel through the incorporation of Fe³⁺ ions, which could interact with both TA and Alg-NHS, leading to the resulting APTF hydrogel patch an “all-in-one” crosslinked network.

The molecular interactions inside the synthesized hydrogel, comprised of multiple macromolecular moieties, are demonstrated in Scheme lb. Initially, PEGDA molecules formed hydrogen bonds with multiple OH groups presented in TA (Schemelb-i). In a previous study, although PEGDA/TA hydrogels were shown to provide improved toughness and wet tissue adhesion in comparison to PEGDA hydrogels, their reported ultimate strengths were below 80 kPa, and the materials demonstrated limited mechanical tunability (6). Additionally, the PEGDA/TA hydrogels lacked multifunctionalities (6). Metal-phenolic networks have recently attracted interest because of their desirable properties (70, 71). In particular, Fe³⁺- phenolic chelation is of particular interest due to its wide broad set of functionalities, including hemostatic, antioxidant and photothermal properties (7, 72, 73). Therefore, it would be highly desirable to incorporate Fe³⁺ into PEGDA/TA adhesive hydrogels to further enhance their properties. However, the facile mono-chelation between TA and Fe³⁺ may interfere with the H-bond interactions between PEGDA and TA, leading to a drop in mechanical properties (Schemelb-ii). To eliminate this trade-off, we incorporated Alg-NHS, which competes with TA for Fe³⁺ ions, in the hydrogel network to increase the dynamic crosslinking densities via ionic crosslinking with Fe³⁺ ions, while also re-introducing the H-bond interactions between TA and PEGDA. Furthermore, synergistic adhesion was achieved through utilizing both TA and NHS to form a wide range of both covalent and noncovalent interactions with the tissues. By controlling the multi-mode molecular interactions, the crosslinking mechanisms in the resulting APTF hydrogel network were modulated, leading to the formation of an ultra-strengthened multi-component adhesive patch (Schemelb-iii).

Synthesis and chemical characterizations of Alg-NHS, AP, and APTF hydrogel.

A schematic for the synthesis of NHS-modified Alg based on the EDC/NHS coupling reaction is presented in Fig. la. The resulting Alg-NHS was chemically characterized with proton nuclear magnetic resonance ('H-NMR), showing the presence of succinimide hydrogen peak at 2.8 ppm (Fig. lb). To increase the NHS content in the Alg backbone, the Alg:EDC:NHS ratio was varied from 1 : 1 : 1 to 1 :2: 10 to 1 :4:20. An increase in the degree of NHS conjugation in Alg-NHS was confirmed by the enhanced NHS peak intensity in the 'H-NMR spectrum of Alg-NHS formed by using higher Alg:EDC:NHS ratio (Fig. lb). Successful conjugation of NHS onto Alg was further analyzed with FTIR spectroscopy, which showed characteristic NHS peaks at 1780 cm'¹ (symmetric carbonyl stretch of NHS), 1704 cm'¹ (asymmetric carbonyl stretch of NHS), 1219 cm'¹ (asymmetric C-N-C stretch of NHS) (Fig. 1c).

The viscosity of the Alg-NHS solution was measured and compared to that of a native Alg solution, as shown in Fig. Id, revealing a lower viscosity for Alg-NHS. Since NHS groups are covalently attached to the carboxyl groups of the Alg moieties which contain two hydrogen bonding, they can restrict the electrostatic interactions of Alg by reducing the negative charge of the carboxyl group. This replaces the electrostatic interactions with dynamic and comparatively weaker hydrogen bonding interactions, reflected in the lowered viscosity of Alg-NHS (2%) compared with that of native Alg (2%) (74).

Both Alg-NHS and TA can interact with Fe³⁺ ions. To understand the effect of Alg-NHS and TA in their cross-interactions with Fe³⁺, rheological studies were performed on Alg-NHS/Fe³⁺ hydrogels and Alg-NHS/TA/Fe³⁺ hydrogels (ATF hydrogels), and the results were compared with those of the Alg/Fe³⁺ hydrogel (Fig. le). While native Alg with bare COOH groups formed a stronger hydrogel with Fe³⁺ ions, the presence of NHS could tune the mechanical properties of the resulting hydrogel by reducing the “egg-box” crosslinking density. Meanwhile, when Fe³⁺ was chelated with TA, it still interacted with Alg-NHS to form a weak hydrogel. This was confirmed upon observing a greater elastic modulus (G’) than the viscous modulus (G”). However, the G’ and G” were lower in the hydrogels containing TA as compared to those without TA. Therefore, we hypothesized that both TA and Alg- NHS may compete for crosslinking with Fe³⁺. Overall, by introducing NHS and TA into the Alg/Fe³⁺ network, the molecular interactions and crosslinking densities across the APTF hydrogels were manipulated to tune their mechanical properties.

As a key functional moiety for tuning tissue-material interfacial interactions (1, 75-77), NHS presence within the AP hydrogel was assessed via FTIR measurements shown in Fig. If. Peaks at 1780 cm'¹ and 1704 cm'¹ in the FTIR spectra of the AP hydrogel (AP hydrogel) are related to the carbonyl stretching of NHS. The peak in 1219 cm’¹ corresponds to the C-N-C stretch of NHS. Formation of APTF hydrogel was also assessed with X-ray photoelectron spectroscopy (XPS) studies, which confirmed the presence of N and Fe elements in the ATPF hydrogel (blue shaded areas in Fig. 1g, which were absent in PEGDA hydrogel). The deconvoluted XPS spectrum of the carbon region showed the presence of 0=C-0 (-288.5 ev, from Alg-NHS), O=C-N (-287 ev, from Alg-NHS) and C-0 (-286.5 ev, from PEGDA) in the hydrogel network. Meanwhile, the presence of TA in the hydrogel was confirmed by pi-pi* interactions visible in the region of 291.8 ev (Fig. Ih). Nitrogen XPS further demonstrated the presence of NHS in both AP hydrogel and APTF hydrogel (Fig. SI in Zheng Bioactive Materials).

Physical properties of APTF hydrogel

Since PEGDA forms the primary skeleton of the APTF hydrogel, it is important to understand the interactions between TA and PEGDA along with the role of Fe³⁺ in controlling the crosslinking of the hydrogel. Fig. 2a shows that the TA- crosslinked PEGDA hydrogel formed without Fe³⁺ underwent nonhomogeneous crosslinking, resulting in swelling in some portions of the PEGDA/TA hydrogels. This could be due to the diffusion-governed crosslinking, during which the surface of PEGDA had a higher crosslinking density than the bulk interior portion, leading to the formation of an inhomogeneous crosslinked network and resulting hydrogel deformation. Interestingly, this effect was eliminated by adding Fe³⁺, which may decrease the crosslinking between TA and PEGDA as it underwent mono-chelation with TA. Indeed, as shown in Fig. S2 in Zheng Bioactive Materials, increasing the concentration of Fe³⁺ led to a decreased amount of TA crosslinked into the hydrogel network due to the blocking of H-bonds between PEGDA and TA. Hydrogels composed of PEGDA/TA/Fe³⁺ are denoted as PTF hydrogels.

Meanwhile, physically encapsulating Alg-NHS inhibited nonhomogeneous crosslinking as Alg-NHS possessed fewer interactions with TA as compared to PEGDA. TA mainly interacted with PEGDA and had minimal interactions with Alg- NHS, which was confirmed by increasing the Alg-NHS concentration in the AP hydrogel (at a fixed total polymer concentration of 22%) leading to a decreased amount of TA incorporated into the hydrogel during crosslinking (Fig. S2 in Zheng Bioactive Materials). By interlocking the PEGDA polymer chains with Alg-NHS, a more homogeneous crosslinking outcome was achieved, as shown in Fig. S3 in Zheng Bioactive Materials.

Since TA was incorporated into the hydrogel through non-covalent interactions, a release study of TA was performed over time. Although there have been numerous studies focused on TA-crosslinked hydrogels, few studies have reported the release kinetics of TA over time (6). TA leaching may be beneficial due to its antibacterial, antioxidant, and anti-inflammatory effects (78, 79). Therefore, TA was used as a releasing agent for accelerating wound healing. For example, Neethu et al. developed a pH-sensitive hydrogel for sustained delivery of TA to promote wound healing (80). However, a localized high TA concentration (>3 mg/ml) could have slight cytotoxic effects (81). As a result, controlling the release of TA is critical for successfully executing its function in a proper manner. Figs. 2b-c; and S4 in Zheng Bioactive Materials demonstrate the release profiles of TA from the engineered hydrogels, formed under different treatments (40% TA/0%, 3%, 6% Fe³⁺) over time. The decreased amount of TA leaching upon Fe³⁺-assisted crosslinking was due to 1) less TA incorporation in the hydrogel as a result of chelation between Fe³⁺ and TA which blocked the H-bonding interaction between TA and PEGDA during crosslinking, and 2) the conversion of the mono-chelated TA/Fe³⁺ within the hydrogel to tris-chelated upon exposure to physiological pH. This pH-sensitive nature of the chelation between TA and Fe³⁺ may serve to stabilize TA within the hydrogel network (75). Therefore, the release profile of the proposed TA/Fe³⁺ crosslinked hydrogel is highly desirable and controllable. The morphologies of different hydrogel compositions were characterized using a scanning electron microscope (SEM). As shown in Fig. 2d, the pure PEGDA hydrogel (Fig. 2d-i) showed the largest pore size compared with the other hydrogel compositions, presumably due to the lower crosslinking density. Interestingly, when the PEGDA hydrogel was interpenetrated with Alg-NHS, the morphology changed completely (Fig. 2d-ii). The effect of Fe³⁺ in contributing to homogenous crosslinking was also manifested in the SEM images, as the PEGDA/TA hydrogel (Fig. 2d-iii) demonstrated a larger pore size distribution compared with that of the PTF hydrogel (Fig. 2d-iv). The final APTF hydrogel demonstrated a more crosslinked structure with the highest roughness and fewest pores when compared to other hydrogels (Fig. 2d-v).

Mechanical characterizations of APTF hydrogels

Controlling intermolecular interactions and crosslinking density within the hydrogel network is the key for tuning its mechanical properties. While TA has been shown to improve the mechanical strength of hydrogels, controlling the mechanical properties of TA containing hydrogels is limited when primarily relying upon Id- bonding interactions (6, 82-86). On the other hand, ion crosslinked Alg hydrogels are known for their brittleness and weak mechanical properties (with a Young’s modulus of 17 kPa and ultimate strain of 20%) (87), significantly limiting their applications due to their cohesive failure during application. To address this limitation, Jeong-Yun et al. prepared a Ca-alginate/polyacrylamide (PAAm) hybrid hydrogel, through forming a ionically and covalently crosslinked network with significantly strengthened toughness as compared to their parents-the alginate and polyacrylamide hydrogels (87). Their work suggested that the fracture energy of resulting hydrogels could be greatly increased by combining weak and strong crosslinks. Here, we hypothesize that by incorporating Alg-NHS/Fe³⁺ into the TA crosslinked PEGDA hydrogel, the mechanical properties would be significantly increased. The tunable properties originate from the Fe³⁺ which controls those two crosslinking systems through different interaction mechanisms with Alg-NHS and TA.

To optimize the APTF hydrogel, we explored the impact of applying different treatment methods on the mechanical properties of the resulting hydrogel. The results highlighted that it is essential to add TA and Fe³⁺ simultaneously to optimize the mechanical properties of the APTF hydrogel (Fig. 3a-d). Treating the AP hydrogel with TA first led to Alg-NHS leaching from the hydrogel network over time. This was confirmed by precipitation of the Alg-NHS in the TA solution after crosslinking using ethanol, whereas alternatively no Alg-NHS precipitated out after co-treating with the TA/Fe³⁺ solution (Fig. S5 in Zheng Bioactive Materials). The released Alg-NHS precipitated out due to a lack of H-bond interactions between Alg-NHS and TA, thereby reducing the crosslinking density between Alg-NHS and Fe³⁺ in the following step. On the other hand, treatment with Fe³⁺ first resulted in a crosslinked Alg- NHS/Fe³⁺ network within the AP hydrogel network. This was confirmed by limited swelling of the AP hydrogel when placed in Fe³⁺ solution as compared to Dulbecco's phosphate-buffered saline (DPBS) after 24 h (Fig. S6 in Zheng Bioactive Materials). However, the limited amount of Fe³⁺ crosslinked with Alg-NHS in the first step would be chelated by TA during the subsequent treatment. Considering the much higher amount of TA compared to Alg-NHS, the crosslinking density between Alg- NHS and Fe³⁺ was lower during the second treatment with TA, leading to weakened mechanical properties. This decreased crosslinking density was confirmed by the darkening of the TA solution after immersing the Alg-NHS/PEGDA/Fe³⁺ (APF) hydrogel into it, indicating chelation of TA and Fe³⁺overtime as the TA/ Fe³⁺ solution is dark in color unlike the clear TA solution featuring with an orange color. As such, we chose to use a TA/Fe³⁺ co-treatment for the APTF hydrogel synthesis, which resulted in the highest Young’s modulus, ultimate strength, and toughness due to the optimal crosslinking between Alg-NHS and Fe³⁺.

Nanoparticle formation was not observed in TA/Fe³⁺ complex as confirmed by transmission electron microscopy (TEM) and dynamic light scattering (DLS) studies on the 40% TA/3% Fe³⁺ solution. No nanoparticles were detected based on TEM images (Fig. S8a in Zheng Bioactive Materials). As shown in the DLS result, the size distribution of 40% TA/3% Fe³⁺ solution showed no difference compared with that of Milli Q water, confirming the absence of any nanoparticles (Fig. S8b in Zheng Bioactive Materials). Additionally, there is a strong H-bonding interaction between TA and PEGDA (6, 90). To investigate the contribution of each polymer (Alg-NHS and PEGDA) to the crosslinking, the total polymer concentration (22% (w/v)) was kept constant, while the ratio of Alg-NHS: PEGDA within the AP hydrogel was varied from 0:22, to 4: 18, and 8: 14, followed by treatment with 40% TA containing either 0, 3 or 6% Fe³⁺. The effect of Fe³⁺ concentration on the three AP hydrogel compositions with different treatments was systematically examined.

As shown in Fig. 4a(i), increasing Fe³⁺ concentration led to a decrease in the Young’s modulus of the PEGDA hydrogel (0:22). This was due to the decrease in H- bonding between PEGDA and TA, since the chelation between TA and Fe³⁺ proved to be the stronger interactions compared with H-bonding (91). This decrease in Young’s modulus was not observed for Alg-NHS/PEGDA at 4: 18 (%) due to the crosslinking established between Alg-NHS and Fe³⁺. This trend was made even more prominent upon further increasing the proportion of Alg-NHS (8: 14). The same behavior was also observed for the ultimate strength and toughness of all hydrogels (Fig. 4a(ii-iii)). In fact, the chelation of Fe³⁺ with TA reduces the H-bonding interaction between TA and PEGDA, thereby lowering the stiffness. On the other hand, the interaction between Fe³⁺ and Alg-NHS can improve the mechanical properties of the resulting hydrogels. Therefore, the positive effects of increasing the Fe³⁺ concentration were more obvious when the proportion of Alg-NHS was also increased in parallel. After substituting PEGDA with Alg-NHS, the primary covalent network density decreased, which explained why the 8: 14 AP hydrogel with TA/ Fe³⁺ treatment demonstrated lower stiffness when compared to the 0:22 PEGDA hydrogel with TA treatment. Meanwhile, simultaneously increasing the Alg-NHS concentration and decreasing PEGDA concertation led to enriched noncovalent interactions, and a resulting improvement in deformability with significantly improved stretchability of up to 1000% (Fig. 4a(iv)).

To understand the net effect of Alg-NHS on the mechanical properties of the hydrogel, different concentrations of Alg-NHS were added to the 22% PEGDA hydrogels which were further treated with 40% TA and 3% Fe³⁺. Improved mechanical properties (increased ultimate strength and Young’s modulus without compromising stretchability) of the hydrogels were observed with increasing Alg- NHS concentrations. This confirmed the enhanced non-covalent molecular interactions in the engineered hydrogel (Fig. 4b). The effect of PEGDA concentration on the mechanical properties of the APTF hydrogel was evaluated at a constant TA (40%), Alg-NHS (8%) and Fe³⁺ (3%) concentration. Herein, increased Young’s modulus and hydrogel strength were attributed to stronger covalent interactions at higher concentration of PEGDA which could not reversibly break and heal. Therefore, due to the lack of contribution from dynamic noncovalent interactions, the stretchability of the hydrogel decreased dramatically as PEDGA concentration increased (Fig. 4c). Moreover, increasing Alg-NHS concentrations could increase the toughness of the hydrogels, which could not be achieved by increasing PEGDA concentrations (Fig. 4d; and S9 in Zheng Bioactive Materials). Overall, an optimal combination of covalent and non-covalent molecular interactions inside the crosslinking network provided the optimal mechanical properties.

After systematic characterization and optimization, the composition of 22% PEGDA, 8% Alg-NHS, crosslinked with 40% TA and 3% Fe³⁺ possessed the best array of desirable and broadly applicable mechanical properties for the APTF hydrogel. In this optimized hydrogel, the Young’s modulus matched well with that of soft tissue (84 kPa), a strain of 924%, an ultimate strength of 951 kPa, and a toughness of 4697 kJ/m³ were obtained, making this composition a suitable biomaterial for sealing elastic tissue (Fig. 4e-g). A representative strain-stress curve of the optimized APTF hydrogel is shown in Fig. 4e. High elasticity of APTF is shown in Fig. 4f-g. We further confirmed that the APTF hydrogel’s strong mechanical properties, including its superb elasticity, toughness, and ultimate strength, could not be obtained in the case of a PEGDA/TA hydrogel, despite incorporating an increased PEGDA concentration (38% PEGDA crosslinked with 40% TA) (Fig. S10 in Zheng Bioactive Materials). As a functional bioadhesive, it is imperative that the hydrogel maintains its structural integrity throughout the duration of application to tissues. Most bioadhesive hydrogels lose their mechanical integrity over a short period under wet conditions due to their excessive swelling (35-39), or degradation (33, 34). The mechanical stability of the APTF hydrogel was assessed in DPBS at 37 °C over time (Fig. Sil in Zheng Bioactive Materials). Although a drop was observed in the mechanical properties of the APTF hydrogel after 96 h, it still exhibited characteristics superior to those of the majority of freshly prepared TA-based bioadhesive hydrogels reported thus far (6, 82- 86), featuring values of 31 kPa ’s modulus, 798 % ultimate strain, 398 kPa ultimate strength, and 1409 kJ/m³ toughness.

Adhesive characterization of APTF hydrogel

The ability to achieve durable and rapid adhesion to the surface of wet tissue is a key capability limiting the success of bioadhesives. The polyphenol-rich Tannic Acid (TA) molecule is an excellent candidate for incorporation into bioadhesives due to its ability to form various interactions (chelation, H-bonding, hydrophobic interactions, Schiff base reaction, Michael type addition) with native tissues through its five-arm structure (33, 75, 83, 84). The presence of catechols in TA, the primary moieties in muscle foot proteins (mfps) of marine mussels to provide wet adhesion underwater, attributes enhanced tissue adherence. Initially, TA interacts with the tissue through H-bonding. Subsequently, covalent interactions with the tissue (i.e., Michael addition, Schiff base reactions) come into play after catechol moieties in TA undergoes auto-oxidization overtime to form quinones (31, 92, 93). These covalent interactions provide durable adhesion of the material to the tissues during their physiological function. An ideal adhesive biomaterial should be functionalized with binding moieties that can spontaneously form both covalent and noncovalent interactions with tissue surfaces while conforming to deformation. Adhesive moieties that have demonstrated high reactivity and spontaneous covalent interactions with tissue surfaces include aldehyde (94-97), A-hydroxysuccinimide (NHS) ester (1, 13, 75-77), isocyanates (98, 99), and aryl azides (100).

We demonstrated that the APTF hydrogel exhibited strong tissue adhesion, in which NHS and TA jointly acted to form robust and instant covalent and non-covalent bonds with tissue surfaces (Fig. 5a). Many studies have shown compromised adhesion due to the pre-oxidation of catechol or pyrogallol prior to application (101-104), as oxidation can trigger the self-polymerization of TA by converting catechols and pyrogallols to highly reactive quinones that can react with the molecule itself rather than the tissue, thus losing the adhesive moieties (105). Catechols start to get irreversibly oxidized, forming quinone structures when the pH is higher than 5.5 with negligible adhesion at a pH of 7.5 (105). Marine mussels themselves have adopted a smart strategy to overcome the drop in their wet surface adhesion caused by catechol oxidation by generating an isolated acidic environment secluded from seawater and secreting the thiol-rich mfp-6 as a reducing agent (106-110). Our system was designed in a similar fashion, as the APTF hydrogel was formed in an acidic pH environment that prevented the catechol/pyrogallol from undergoing oxidation, thus preserving its strong adhesiveness. American Society for Testing and Materials (ASTM) standard lap shear tests (ASTM F2255) were conducted on freshly cut porcine skin tissue in order to quantify the adhesion efficacy of the engineered APTF hydrogel patch. At a fixed total polymer concentration (22%) with co-treatment of 40% TA and 3% Fe³⁺, increasing Alg-NHS concentration had no effect on the adhesive strength. This was due to lower TA incorporation when less PEGDA was used (Fig. S12 in Zheng Bioactive Materials).

While previously reported PEGDA hydrogels showed no adherence to tissue surfaces due to their nonadhesive nature toward cells and proteins (111), treatment of PEGDA with TA enhanced the adhesive strength of the PEGDA/TA hydrogel to 32.5 ± 2.05 kPa. Herein, TA molecules with multiple hydroxyl groups interact with the polar functionalities of the tissue surface to facilitate adhesion. The contribution of Alg-NHS to tissue adherence was demonstrated by comparing PEGDA and AP hydrogels following TA treatment. The presence of Alg-NHS increased the adhesion strength significantly from 32.5 ± 2.05 kPa for PEGDA/TA hydrogel to 52.24 ±3.86 kPa for the APT hydrogel due to the introduction of NHS-mediated covalent and noncovalent interactions (Fig. 5b). Meanwhile, the incorporation of Fe³⁺ into the AP hydrogel, which crosslinked the Alg-NHS moieties within the hydrogel, further improved the adhesive strength to 77.28 kPa ± 3.03 kPa by preventing the leaching of Alg-NHS from the hydrogel during crosslinking. Fig. 5c represents the shear strengthdisplacement curves related to the engineered hydrogels, reflecting the origin of adhesion from TA and Alg-NHS. Additionally, we found that some NHS groups could undergo acid-triggered ester hydrolysis during crosslinking, which is a reversible (112) and temperature dependent reaction (113). The hydrolysis of NHS in the TA solution was characterized by 'H NMR (Fig. S13 in Zheng Bioactive Materials) which showed that part of the NHS ester was hydrolyzed during the crosslinking. This opens up possibilities for optimizing the solvent and temperature during crosslinking in the future. Quick adherence after 5 sec of gentle pressing on different types of wet tissues (e.g., liver, lung, stomach, heart, muscle, and others) is assessed with the engineered APTF hydrogel and presented in Fig. 5d. One side of the patch was adhered to a spatula, while the other was adhered to the wet tissue (Video SI in Zheng Bioactive Materials). To evaluate the bonding effect of the APTF hydrogel, we further characterized the adhesiveness of the APTF hydrogel through a tensile pull-off test (ASTM F2258). As shown in Fig. 5e, the APTF hydrogel exhibited an adhesive strength of 77.28±3.39 kPa based on the lap shear test and 102.8±3.39 kPa based on the tensile pull-off test, which were significantly higher than that of commercial glues including Coseal, DuraSeal, Histoacryl, BioGlue, and Tisseel (less than 40 kPa) (1). Additionally, the adhesive strength of APTF hydrogel was also higher than most of the recently developed bioadhesives (40, 58, 114, 115). The APTF bioadhesive also exhibited strong adhesion to a wide range of wet tissue surfaces including stomach (13.72=1=2.81 kPa), cornea (19.61±3.26 kPa), and conjunctiva (11.87±2.57 kPa) (Fig. 5f). Removal of patches from wounds is difficult due to the risk of secondary injury and bleeding (116). Especially for infants and diabetic patients with fragile and sensitive skin, extra-strong adhesion might exacerbate pain and cause trauma and inflammation upon removal (117). In addition, on-demand detachment of bioadhesives is crucial for repositioning misplaced bioadhesives or for retrieving implants (52). Achieving tough adhesion and benign triggerable detachment remains a challenge. For instance, many recently developed bioadhesives required harsh triggering conditions which are incompatible for the adjacent native tissues such as concentrated metallic ions, pH, UV irradiation, or electrical stimulus to induce the detachment (118-121). Recently, Jiang et al. developed a gelatin-based bioadhesive that could be removed by applying cold water. However, their hydrogels suffered from low adhesive strength (less than 10 kPa) (65). Since APTF hydrogel possesses strong adhesive properties, therefore, endowing APTF hydrogel with a chemically tunable debonding characteristic is beneficial. Herein, the debonding capability of the engineered APTF patch was demonstrated in Fig. 5g-h which reveals facile patch removal upon the application of deferoxamine mesylate (DFO) and urea solutions at different concentrations. As a strong medical Fe³⁺ chelator, DFO is FDA approved and widely used in clinical cases of acute iron intoxication (122). It has been shown that DFO readily interacted with Fe³⁺ ions involved in the crosslinking of hydrogels, leading to the resulting dissociation of the hydrogel (7). Therefore, DFO is expected to reduce the adhesion of the hydrogel by decrosslinking and subsequently decomposing the hydrogel network. Meanwhile, urea is a naturally occurring molecule, which is produced during protein metabolism and present in large amounts in human blood and urine (123, 124). Urea molecules have a superior tendency to form hydrogen bonds and can thus interact with interfacial TA molecules to weaken the adhesion of the hydrogel patch to the tissue surface (6). Here, we showed that using a higher concentration of urea significantly lowered the adhesion strength of the patch (Fig. 5h). The on-demand detachment characteristics were also assessed using wet tissues such as the cornea and conjunctiva. As is shown in Fig. 5i, the application of 50 pL urea (0.1 M) could significantly reduce the adhesion to the tissue substrates. As the required concentrations and amounts of the urea and DFO for detachment were remarkably lower than their safe dosage used in clinics (124-128), we do not expect to see any side effects with using them.

To demonstrate handling and clinical application of the designed material, tissue adhesion was tested in situ using a rabbit eye conjunctiva and pig cornea. Cornea and conjunctiva were adopted to assess the wet adhesion of the designed APTF hydrogel since these tissues are known for their consistent secretion of tear fluid and mucin, which lubricate the eye but limit surface adhesion (129). To the best of our knowledge, there is no ready-to-use bioadhesive patches developed for quick adherence to the eye due to the slippery and extremely wet surface. To demonstrate the potential in vivo wet adhesion of the APTF bioadhesive to the cornea and conjunctiva, adhesion was assessed immediately after euthanizing the rabbit and pig when the cornea and conjunctiva surfaces were fresh and wet. We demonstrated that the APTF patch strongly adhered to the eye conjunctiva upon gentle pressing for one min, and the wet adhesion was immediately assessed by attempting to remove the patch with a tweezer (Fig. 5j). Further, in addition to the demonstration of robust adhesion formation, the painless on-demand detachment of the APTF hydrogel was also shown by applying 2-3 drops of a 0.1 M urea solution to the adhesion site on both animal models (Video S2 and S3 in Zheng Bioactive Materials).

In vitro biocompatibility assay of APTF hydrogel

To evaluate the biocompatibility of the APTF hydrogel, 3T3 cells were drop- seeded on top of the AP and APTF hydrogels, followed by incubation for 7 days. NIH 3T3 fibroblasts were selected because it is one of the most frequently used cell lines for studying material biocompatibility and cytotoxicity in accordance with the International Organization for Standardization (ISO) norm 10993-5 (130-133). A live/dead assay demonstrated excellent cellular viability (> 90%) for both AP and APTF hydrogels up to 7 days post-seeding (Fig. 6a-b, S14 in Zheng Bioactive Materials). Fluorescent F-actin/cell nuclei staining was also performed to demonstrate cell proliferation and spreading on hydrogels up to 7 days post-seeding (Fig. S15). However, due to the lack of arginine-glycine-aspartic acid (RGD) moieties in the hydrogel, the cells proliferated relatively slowly on both the AP and APTF hydrogels as demonstrated by a PrestoBlue assay (Fig. 6c). To further evaluate the biocompatibility of the APTF hydrogel, we cultured NIH 3T3 cells in Transwell® cell culture inserts and exposed them to APTF hydrogel. In this method, NIH 3T3 cells were seeded at the bottom of a 24-wells Transwell permeable (Costar®, 8 pm PET membrane) at 2* 10⁴ cells/cm². APTF hydrogels were placed into Transwell inserts, and 1 mL of growth medium (Dulbecco's Modified Eagle's Medium) was added to each well of the Transwell permeable supports. As shown in Fig. S16 in Zheng Bioactive Materials, based on a live/dead assay, cells treated with APTF hydrogel showed no difference compared with that of the control group, confirming the in vitro biocompatibility of the APTF hydrogel.

Antioxidant activities of APTF hydrogel

Free radicals generated at the wound site may cause oxidative stress and cytotoxicity by damaging DNA and enzymes (134). Antioxidant bioadhesive patches can facilitate the wound healing process by scavenging excess reactive oxygen species (ROS). The antioxidant properties of the APTF patches were determined by evaluating the scavenging ability of the hydrogels against the stable free radical colorimetric probe, l,l-diphenyl-2-picrylhydrazine (DPPH»), where the color change can be quantified by UV-vis spectrophotometry (135). TA is known to have antimutagenic and antioxidant activities. The antioxidant activity of TA is attributed to its capacity to form a complex with ferric ions, interfering with the Fenton reaction (136). Here, to evaluate the antioxidant properties of the engineered hydrogels, the color changes of the three DPPH* solutions exposed to hydrogel-lacking control, the AP hydrogel and the APTF hydrogel were monitored at different time points. As shown in Fig. 6d, the color of the DPPH* solution exposed to the APTF hydrogel changed from dark purple to light yellow within 10 min, which was not observed in DPPH* solutions exposed to either the hydrogel-lacking control or the AP hydrogel. Typically, DPPH radicals gave a strong absorption at 517 nm in the UV visible spectroscopy. After being reduced, the strong peak at 517 nm disappeared only in the case of the APTF hydrogel (Fig. 6e). We found that the APTF hydrogel possessed strong antioxidant activities with 91% radical scavenging efficiency due to the presence of TA/Fe³⁺ complexes (Fig. 6f).

Photothermal properties and Near-infrared (NIR)-assisted antimicrobial activity of APTF hydrogel.

Photothermal ablation, which utilizes photothermal agents under NIR radiation (700 nm-1400 nm), has been applied to kill antibiotic-resistant bacteria in the treatment of infections (137). Consequently, mild photothermal ablation has received significant attention as an alternative treatment to conventional antibiotics due to its remote deep tissue penetration to kill bacteria without damaging normal tissues. On the other hand, compared to traditional photothermal agents such as Au nanoparticles, the hydrogel compound consisting of catechol-Fe³⁺ can protect normal tissues from thermal damage while mitigating adverse effects such as nanomaterial diffusion (71).

Therefore, we envision that the APTF hydrogel containing TA/Fe³⁺ can effectively absorb and convert NIR light to heat for antibacterial applications. To test this, an 808nm NIR laser was used to characterize the photothermal properties of the APTF hydrogel under wet and dry conditions. After irradiation for 4 min, the temperature increments of a dry APTF hydrogel were 27.6 °C and 38.6 °C at 0.015 W/cm² and 0.016 W/cm², respectively (Fig. 7a). For a wet APTF hydrogel, the temperature dropped to 22.53 °C at 0.015 W/cm², and 37.8°C at 0.016 W/cm² (Fig. 7b). However, the pure AP hydrogel showed almost no temperature increase at 0.016 W/cm² in either dry or wet conditions (~5°C and ~ 0°C), indicating a lack of photothermal effects. Meanwhile, pyrogallol-Fe³⁺ crosslinked APTF displayed a tunable photothermal capacity, depending on the intensity of NIR light exposed to the hydrogel surface. A cyclic photothermal heating test demonstrated that the APTF hydrogel could be heated to 50 °C in 2 min during irradiation. Upon removal of irradiation, the APTF hydrogel could cool down to room temperature within 2 min. The application of 10 consecutive heating cycles demonstrated the photothermal stability of the APTF hydrogel, indicating no long-term degradative photothermal effects (Fig. 7c). Moreover, the infrared images showed that the maximum photothermal temperature of the APTF hydrogels was quickly reached in 1 min and remained stable over 5 min, confirming the practical usability of the hydrogel in both dry and wet conditions (Fig. 7d).

Our APTF hydrogel demonstrated NIR-assisted antibacterial functionality because the hydrogel reached and exceeded a photothermal temperature of 45 °C, beyond which the viability of some bacterial enzymes begins to rapidly decline due to denaturation, resulting in bacterial death (7, 138). The NIR-assisted antibacterial property of APTF hydrogels was then assessed against Pseudomonas aeruginosa (P. Aeruginosa) and Methicillin-resistant Staphylococcus aureus (MRSA). After irradiation for 5 min, each survived bacteria sample was seeded with a 1 mL aliquot of DPBS solution. After an additional 5 min, bacteria samples were spread onto agar plate, incubated overnight, and the colonies were counted. Our results revealed nearly complete eradication of P. Aeruginosa when the NIR assisted APTF hydrogel was used with a log reduction of 5.389. Meanwhile, NIR light alone provided no bactericidal effects. Among the tested samples, only the combination of the APTF hydrogel assisted with NIR light interacted synergistically to demonstrate antimicrobial activity whereas the APTF hydrogel without NIR light assistance showed no difference in bacterial viability compared to the control (Fig. 7e and g). In addition, the AP hydrogel with NIR light also showed zero antibacterial activity. The strong antibacterial properties of APTF hydrogel were confirmed against both gram positive bacteria (MRSA) with only 8.733 ± 7.12 % bacterial viability and gram negative (P. Aeruginosa) bacteria with complete eradication (Fig. 7f), presumably because Staphylococcus aureus can tolerate higher temperatures when compared with other bacteria (139). Having potent effects for killing bacterial without causing cellular toxicity is important for the safe usage of APTF bioadhesive for medical applications. Compared with the control group (cells without any treatment), the APTF hydrogel coupled with NIR had a mild cytotoxic effect on 3T3 cell viability (Fig. 7h and S17). In general, sustained exposure to high temperatures (above 50 °C ) may cause irreversible damage to normal cells and tissues (140, 141), while heat- induced cell death (e.g., apoptosis) at lower temperatures (42-47 °C) can be reversed with the help of heat shock proteins (142, 143). The optimal temperature for enzymatic activity in bacteria is between 30 and 40 °C and is strongly inhibited at higher temperatures (7, 144-146). Therefore, careful control of irradiation intensity and time is crucial to avoid the side effects of NIR laser on normal cells and tissues while killing bacteria. In one study, Lee et al. developed a hydrogel platform that was photopolymerized through NIR light irradiation for cell delivery. After a 5 min incubation at 45 °C, 3T3 cell viability around 86% was observed, which was comparable to the cells incubated at 37 °C (147). In fact, many studies have shown that a temperature of 40-50 °C can effectively kill bacteria without causing any damage to the cells (146, 148). When the temperature was higher than 50 °C, strong bacterial killing effects were observed (>90% killing efficiency) with mild or partial cell damage, and these cells could rapidly proliferate to a normal state (149-151). In another study, Li et al. developed a TA/Fe³⁺-based hydrogel for NIR-assisted antibacterial wound treatment. The hydrogel could reach a temperature of more than 60 °C that efficiently killed S. aureus and E. coll (> 90% killing efficiency). A certain extent of apoptosis of the fibroblasts was detected. However, fibroblasts around the treated site could rapidly proliferate and reach almost the same number as the control group after irradiation (152).

In our study, we used a FLIR thermal camera to record the temperature of both sides of the hydrogel during application. We found that although the surface of the hydrogel (~ 1mm thickness) facing the light could reach a temperature of 60°C after 5 min irradiation, the back side of the hydrogel, interfacing directly with cells, could achieve slightly above 40 °C, ensuring its safe usage (Fig. S18 in Zheng Bioactive Materials). In a recent work where catechol-Fe³⁺ was adopted for NIR-assisted antibacterial purposes, it was reported that the wound site temperature of the hydrogel group could increase to above 50 °C within 3 min in an in vivo infected skin wound, while the tissue around the hydrogel just showed a slight temperature increase to about 38 °C, and the temperature of other parts of the mouse had no significant change (153). Due to its minimal interaction with living tissue and ability to kill bacteria without causing serious damage to cells, NIR light has been combined with various biomaterials and used for different medical applications. This was demonstrated in a variety of reported in vivo wound healing models, where NIR- assisted therapy showed the best wound healing outcomes with the least inflammation, best regenerative results, and no tissue damage compared to other groups with no NIR treatment (7, 146, 148, 150-152, 154-161). Our results together suggest that APTF hydrogel can provide antibacterial effects with no significant damage on cells.

We have demonstrated the development of a multifunctional bioadhesive with selectively tunable physical properties for the sealing and treatment of injured tissues. The engineered bioadhesive hydrogel was formed via a combination of Alg-NHS, PEGDA, TA and Fe^{J ;}. By fine-tuning the covalent and noncovalent molecular interactions within the hydrogel network, the resulting APTF hydrogel demonstrated excellent toughness, appropriate stiffness, and improved ultimate strength and elasticity without any tradeoffs. Long-term mechanical integrity was also well- preserved under wet conditions. In addition, robust and instant adhesion was achieved through the synergistic effects of dual adhesive moieties (NHS and TA) in the hydrogel, governing the tissue-material interfacial interactions. We also demonstrated the on-demand removal of the designed adhesive patch through the direct application of nontoxic agents to the tissue-material interfaces. Meanwhile, the presence of TA/Fe³⁴ imparted multiple functionalities to the hydrogel system, including antibacterial and antioxidant properties suitable for wound healing. Additionally, the hydrogel demonstrated excellent in vitro biocompatibility. This work showcases the tunability of a single hydrogel platform with unique physio-chemical characteristics, including superior mechanical, adhesive properties, along with multifunctionalities (i.e., antioxidant and antibacterial effects), making it suitable for the clinical treatment of wounds across diverse tissue types. Furthermore, this biomaterial platform gives way to the prospect of designing and molecularly engineering a class of tissuespecific biomaterials for various biomedical applications, such as internal and external hemorrhage control, drug delivery patches, and others. Our future work will focus on assessing the in vivo biocompatibility and wound healing effects of the engineered APTF hydrogel using different animal models to broaden its application as bioadhesives for wound sealing and repair, hemostatic patches for hemorrhage control as well as matrices for drug delivery.

Materials and Methods

Materials

Poly (ethylene glycol) (PEG), alginic acid sodium salt, acryloyl chloride, hydroquinone, tannic acid (TA), urea and lithium phenyl-2,4,6- trimethylbenzoylphosphinate (LAP) were all purchased from Sigma-Aldrich. Deferoxamine mesylate was purchased from ApexBio. l-(3-Dimethylaminopropyl)-3- ethylcarbodiimide Hydrochloride (EDC HC1), A-Hydroxy succinimide (NHS) and 1,1- Diphenyl-2-picrylhydrazyl Free Radical (DPPH) were obtained from TCI chemicals. Iron (III) chloride, anhydrous was purchased from Alfa Aesar. Dulbecco’s modified Eagle medium (DMEM) was purchased from Cellgro (Manassas, VA), Fetal Bovine Serum (FBS) and Dulbecco’s phosphate-buffered saline (DPBS) were obtained from HyClone (Logan, UT). Live/Dead viability kit, and penicillin/streptomycin (Pen- Strep) were purchased from Invitrogen (San Diego, CA).

Synthesis of PEGDA PEDGA was synthesized through chemical conjugation of PEG molecules and acryloyl chloride in toluene. First, 30 g of PEG (20 kDa) was dissolved in toluene and mixed with 2.5 mL trimethylamine. Next, a solution containing 1 mL of acryloyl chloride and 10 mL of dried toluene was added to the reaction mixture and allowed to react for 2 h. Finally, the reaction mixture was filtered through a silica bed and collected in a flask containing 200 pL of 30-50 ppm hydroquinone solution in acetone. Proton nuclear magnetic resonance (^XH NMR) was used to verify diacrylation, with efficiency calculated at 93%. The chemically modified PEG was precipitated in hexane and stored at -80°C for future use.

Synthesis of Alg-NHS

Alginate (Alg)-NHS was prepared by esterifying the -COOH group in alginic acid sodium salt. One gram of alginate was dissolved in distilled water, followed by the addition of EDC and NHS in different molar ratios. The mixture was allowed to react for 3 h at 45°C, and the final product was precipitated out via addition of cold ethanol, and then washed completely with pure ethanol, and then lyophilized.

APTF hydrogel synthesis

First, a solution of LAP photoinitiator was prepared at a concentration of 0.5 mg/mL in Milli Q water. Alg-NHS and PEGDA were weighed and added to the LAP solution and fully dissolved. After dissolving, the solution was slowly casted in poly(dimethylsiloxane) (PDMS) molds and exposed to UV light (405 nm) for 4 min to polymerize the solution and form a hydrogel. Next, 40% TA and Fe³⁺ solution at varied concentrations (0%, 3%, 6%) were prepared in Milli Q water. The Alg-NHS and PEGDA hydrogels (AP hydrogel) were then immersed in the TA/Fe³⁺ solution and crosslinked for 24 h. Finally, the hydrogels were taken out, washed thoroughly with Milli Q water, and dried in a vacuum overnight. The weight ratio of TA in the composite hydrogels (W%TA) was calculated by the weight loss method:

where Wdry is the weight of dry scaffold of hydrogels. For calculating the weight ratio, the hydrogels were further freeze dried for 5 days.

Hydrogels containing Alg-NHS/PEGDA were denoted as AP hydrogels. Hydrogels containing Alg-NHS/TA/Fe³⁺ were abbreviated as ATF hydrogels. Hydrogels containing Alg-NHS/PEGDA/TA were abbreviated as APT hydrogels. Hydrogels containing Alg-NHS/PEGDA/ Fe³⁺ were designated as APF hydrogels. Hydrogels containing PEGDA/TA/Fe³⁺ stands for PTF hydrogels. Hydrogels containing Alg-NHS/PEGDA/TA/Fe³⁺ were named as APTF hydrogels.

Chemical characterizations of Alg-NHS, AP hydrogel and TA/Fe³⁺.

’II NMR analysis

Chemical functionalization of Alg was confirmed with ¹ H NMR spectroscopy analysis using a 400 MHz Bruker AV400 spectrometer. The samples were prepared by dissolving 10 mg of the dried polymers in 1 mL of D2O. The hydrolysis of NHS of Alg-NHS in TA solution was examined with ¹ H NMR.

FTIR analysis

For the Fourier-transform infrared spectroscopy (FTIR) analysis, 3-5 mg of the samples was grinded with dry KBr to form pellets and saved for the measurements in a PerkinElmer Paragon 1000 FT-IR Spectrometer. The chemical composition of AP hydrogel was analyzed through ATR-FTIR.

Ultraviolet-visible (UV-vis) analysis

The spectrum of the TA/ Fe³⁺ aqueous solution was recorded using a Thermo Scientific NanoDrop One/OneC Spectrophotometer at 200-1000 nm.

XPS analysis An X-ray photoelectron spectrometer (AXIS Ultra DLD instrument) was used to analyze the chemical composition of the PEGDA and APTF hydrogels. A monochromatic Al Ka X-ray at 15 kV and 10 mA was used as an excitation source. The neutral C is peak (C-C (H) set at 284.6 eV) was used as a reference for charge correction.

Rheological characterizations

Rheological investigations were performed by Modular Compact Rheometer MCR302 to characterize the interm olecular interactions. The results were obtained by linking the measuring system PP08 with diameter of 8 mm to the rheometer. Each measurement was carried out by loading a fresh sample in the 1mm gap between the parallel plates and removal of excessive sample. An oscillatory frequency sweep, with 120 measuring points was carried out in the frequency range of 0.1 to 120 rad s ' at room temperature (25°C). At a given shear rate parameter, ranging from 1 to 100 s ' with 120 measuring points, the relationship of viscosity and shear stress as a function of shear rate was recorded. The storage modulus (G’) and loss modulus (G”) were recorded in the linear viscoelastic regime at a shear strain of y = 1% and an angular frequency sweep in the range of 0.1 to 120 rad/s.

Scanning Electron Microscopy (SEM) Analysis

The cross-sections of PEGDA, PEGDA/TA, AP, PTF and APTF hydrogels were sputter coated with a very thin layer of gold-palladium (Au-Pd) alloy for 30 s and examined by SEM (ZEISS Supra 40 VP SEM).

Transmission electron microscopy (TEM) Analysis

A 3 pL of 40%TA/3% Fe³⁺ solution with 10 times dilution in Milli-Q water was added onto a TEM grid (Electron Microscopy Sciences, Formvar/Carbon 200 Mesh, Copper), dried overnight and imaged using a T12 Quick room temperature TEM with 120 kV electron-beam energy. Dynamic light scattering (DLS)

A 40%TA/3% Fe³⁺ solution with 10 times dilution in Milli-Q water was used for DLS analysis using a Malvern Panalytical DLS Zetasizer to confirm particle formation, and Milli-Q water was used as a control.

In vitro mechanic characterizations

To characterize the mechanical properties of the swollen hydrogel, the hydrogel samples were equilibrated in DPBS at 37°C for 1 h before testing. The tensile properties of the APTF patches were measured by conducting pure-shear tensile tests on thin rectangular samples (10 mm in length, 3 mm in width, 0.9 mm in thickness with the precise dimension of each sample measured using a digital caliper) with a mechanical testing machine (Instron 5943). All tests were conducted with a constant tensile speed of 50 mm/min. The strain and ultimate strength were defined as strain and stress at the failure point, respectively. The Young's modulus was calculated from the slope of 1/6 - 1/5 of the strain in the stress-strain curve for all the samples while the toughness was calculated as the area under the stress-strain curve.

Tannic acid releasing study

PEGDA hydrogel was crosslinked with 40% TA and varied concentrations of Fe³⁺ (0%, 3% ,6 %) for 24 h. After crosslinking, the hydrogels were thoroughly washed with Milli Q water, and a release study was carried out by immersing the hydrogels into 5 mL of DPBS and then into a 37 °C oven. After each predetermined time point, 1 mL of the released tannic acid solution was taken out and promptly replenished by 1 mL of fresh DPBS. The amount of TA released over time was quantified through a UV-vis spectrophotometer (NanoDrop One/OneC) at 259 nm and the concentration of TA released over time was determined by a concentrationabsorbance standard curve. In vitro adhesive characterizations

For demonstrative purposes, one side of the APTF patch was attached to a spatula while the other side was firmly adhered to different tissues (lung, stomach, aorta root, heart, intestine, muscle, skin, liver) after 5 sec of gentle pressing.

For the adhesive lap shear test (ASTM F2255) and the tensile pull-off test (ASTM F2258), fresh porcine skin tissue surfaces were completely wetted with DPBS before applying the adhesive patch. After adhering both sides of the patch to the tissues, it was pressed for 1 min for adhesion enhancement before being placed in a 37 °C oven for 15 min to equilibrate and simulate human body temperatures. Lap shear test was performed on the other tissues including fresh porcine stomach, cornea, and conjunctiva using the same procedure described before.

To measure the adhesive strength of the samples, a mechanical testing machine (Instron 5943) was used to conduct a standard lap-shear and tensile pull-off test on the adhered samples (18 mm x 7.5 mm, the precise dimension of each sample was measured using a digital caliper). All tests were conducted with a constant tensile speed of 50 mm/min and the adhesive strength was determined via dividing the maximum force by the adhesion area. For the lap shear test, glass slides layered with cyanoacrylate glue were used to stiffly set the tissues for adhesion to the hydrogels. For the tensile pull-off test, two exterior sides of the porcine tissues, with APTF hydrogel adhered in between, were attached to compression platens using cyanoacrylate glue for pulling off. To assess the detachment of the APTF hydrogel in vitro, 50 pL DFO (10%) and Urea (0.1M and 0.5M) at different concentrations were applied to tissues (porcine skin, cornea and conjunctiva) and subsequently followed by a lap shear test using the previously mentioned procedure.

Adhesion characterizations on rabbit conjunctiva and pig cornea

Adhesion to conjunctiva tissue was evaluated on a New Zealand rabbit and a Yucatan pig model. Adhesion was assessed immediately after euthanizing the rabbit and pig to ensure the freshness of the tissue. Specifically, APTF patch was applied to the conjunctiva tissue of the rabbit or the cornea tissue of the pig and was gently pressed for one min to enhance the adhesion. Upon removing the pressure, the adherence to the tissue was assessed by removing the APTF patch with a tweezer. On- demand detachment was also evaluated in both models by applying 2-3 drops of a 0.1 M urea solution to the adhered patch-tissue interface. Afterwards, the adherence to the tissue was assessed again by removing the APTF patch with a tweezer.

In vitro biocompatibility test

To evaluate the biocompatibility of the APTF hydrogel, a 30 pL AP hydrogel precursor solution was placed into a 150 pm spacer covered by a 3 -(trimethoxy silyl) propyl methacrylate (TMSPMA) coated glass. Afterwards, the AP hydrogel precursor was photopolymerized for 4 min using UV light (405 nm) and then crosslinked with TA/Fe³⁺ for 24 h. After fixing APTF hydrogel to the coated glass slides, 2 x 10⁴ cells/samples were seeded on top of the APTF hydrogel and placed in 24-well plates with 1 mL of growth medium (DMEM + 10% FBS+ 1% Pen- Strep) having pure AP hydrogel as the control group. The 3D cultures were incubated at 37 °C in a 5% CO2 humidified atmosphere for 7 days with the growth medium being replaced every 2 days. Cell viability was determined using a calcein AM/ethidium homodimer- 1 live/dead kit (Invitrogen) according to manufacturer protocol. The viability of 3D encapsulated NIH 3T3 fibroblasts in APTF hydrogel were evaluated on days 1, 5, and 7 post-encapsulations. Fluorescent images were taken using a Zeiss Axio Observer Z1 inverted microscope analyzed with ImageJ software. Cell viability was calculated as the ratio of viable cells to the total number of live and dead cells. The PrestoBlue assay was used to evaluate the metabolic activity of 3D encapsulated NIH 3T3 fibroblasts inside the APTF and AP hydrogels on days 1, 5, and 7 post-incubations. The fluorescence intensity of the resulting solutions was recorded at 535-560 nm excitation and 590-615 nm emission. The spreading of 3T3 cells on APTF hydrogel was visualized through the fluorescent staining of F-actin and cell nuclei. Briefly, cell cultures at APTF hydrogel at days 1, 5 and 7 post-seeding were fixed in 4% (v/v) paraformaldehyde (Sigma) for 15 min, permeabilized in 0.1% wt/v Triton X-100 (Sigma) for 5 min, and then blocked in 1% wt/v bovine serum albumin (BSA, Sigma) for 30 min. Samples were then incubated with Alexa fluor 594 phalloidin for 45 min. After consecutive washes with DPBS, samples were counterstained with 1 pL/mL DAPI (4’,6-diamidino-2-phenylindole, Sigma) in DPBS for 2 min. Fluorescent imaging was carried out using an inverted fluorescence microscope (Zeiss Axio Observer Z7).

The cytotoxicity of the engineered APTF hydrogel was also evaluated using Corning® Costar® Transwell® cell culture inserts. Commercial live/dead kits (Invitrogen) were used to assess the cell viability. APTF hydrogels, prepared following the previously described protocol, were placed inside the Transwell inserts, and 1 mL of growth medium (Dulbecco's modified Eagle's medium) was added to each well of the Transwell permeable support. The well plates were maintained at 37 °C in a 5% humidity environment for 5 days with medium changes every 48 h. The viability of 3T3 cells grown at the bottom of the well plate was assessed using a live/dead cell viability kit. Cells seeded on well plates were stained with 0.5 pL/mL calcein AM and 2 pL/mL ethidium homodimer- 1 (EthD-1) in DPBS for 20 min at 37 °C. Fluorescent imaging was performed with an AxioObserver Z7 inverted microscope on days 1 and 5 post-seeding. Live and dead cells were visualized as green and red color, respectively.

Photothermal characterization of APTF hydrogel

An 808nm Near-infrared (NIR) laser was used to characterize the photothermal properties of the prepared APTF hydrogel under wet and dry conditions, AP hydrogel was used as a control. Wet and dry hydrogels were irradiated for 6 min at a density of 0.015 W/cm² and 0.016 W/cm². Temperature recordings were taken from a Forward-looking infrared (FLIR) thermal camera at 10 sec, 30 sec, and 1-min intervals from the start of the heating time. In addition to the temperature recordings, an infrared picture was taken to create the thermal imaging table for the tested hydrogels. A cyclic test was performed on dry APTF hydrogel by constantly heating the hydrogel for 2 min using an 808 nm laser at 0.015 W/cm², allowing it to cool for 2 min, and reheating the hydrogel shortly thereafter over a span of 40 min.

NIR-assisted antibacterial activity assay

Methicillin-resistant Staphylococcus aureus (MRSA) and Pseudomona Aeruginosa (P. Aeruginosa) were employed to test the antibacterial activity of the APTF hydrogel. For each of the bacteria, three trials were conducted for five different types of samples including sterilized DPBS solution with NIR light, DPBS solution without NIR light, AP hydrogel with NIR light, APTF hydrogel without NIR light, and APTF hydrogel with NIR light. Each hydrogel (67 pL) was placed into a 48 well plate with their surface seeded with 10 pL of bacteria at a concentration 0.12 at optical density of 600 (0.245* 10⁸ CFU/mL for P. aeruginosa (1) and 0.18 * 10⁹ CFU/mL for MRSA (2)) and irradiated at a density of 0.016 W/cm² with an 808nm laser. Well plates which did not contain the hydrogel were instead seeded with 67 pL of sterilized DPBS solution and then seeded with 10 pL of bacteria. Then, each well plate marked to receive NIR light was irradiated for 5 min with the light shining onto the bacteria containing surface of the hydrogel. Afterwards 1 mL of DPBS was placed into each well to resuspend any bacterial survivors, allowed to mix for 5 min, and then 1 mL of the suspended solution was spread onto agar plate and incubated for 24 h at 37.0 °C. Afterwards macroscopic pictures and colony counts were performed on each of the agar plate using a Zeiss microscope. Bacterial viability was calculated through dividing colony counts of experimental group by control group. Antibacterial efficiency was also expressed using log reduction which is equal to the Log cell count of control minus Log survivor count on hydrogel after NIR-assisted antibacterial treatment (3).

NIR-assisted cytotoxicity assay To evaluate the cytotoxicity of NIR generated heat during photoablation, NIH 3T3 fibroblasts were seeded on 48 well plate (2 x io⁴ cells). APTF patch was placed on top, followed by 808 nm irradiation for 5 min at 0.016 W/cm², and then 5 min equilibrium time. An infrared picture was taken on the front side (facing irradiation) and on the back side (facing cells) of the hydrogel using a FLIR thermal camera to record the temperature immediately after irradiation. Next, APTF hydrogels were taken out, and cell viability was determined using a calcein AM/ethidium homodimer- 1 live/dead kit (Invitrogen) according to manufacturer protocol. Fluorescent images were taken using a Zeiss Axio Observer Z1 inverted microscope analyzed with ImageJ software. Cell viability was calculated as the ratio of viable cells to the total number of live and dead cells.

Antioxidant activity characterizations of APTF hydrogel

The antioxidant efficiency of the prepared hydrogels was evaluated by the reported method which scavenge the stable 1, l-diphenyl-2-picrylhydrazyl (DPPH) free radical (4). Specifically, DPPH was dissolved in the ethanol to form the solution with certain concentration (0.2 mmol/L); 200 uL APTF hydrogel and AP hydrogel were mixed with the 3 ml of the DPPH/ethanol solution. The mixture was placed in a vibrating shaker at a speed of 100 rpm, and incubated in the dark for 10 min. Then, the wavelength scanning and absorbance at 517 nm of the solution was obtained by a UV-vis spectrophotometer (NanoDrop One/OneC Spectrophotometer). The DPPH radical scavenging rate was calculated by the following formula:

DPPH scavenging %=(Ao-Ai)/Ao^x lOO % where Ao was the absorption of the DPPH solution, and Ai was the absorption of the DPPH solution after reacting with the hydrogel samples.

Statistical analysis The statistical analysis method (One-way ANOVA) was used to assess the significance of the assay data, p < 0.05 was indicated as the level of significance. When the data has p < 0.05, it is indicated with (*), and (**) means that p < 0.01, (***) indicates that p < 0.001, and (****) specifies that p <0.0001. For every experiment, at least three samples were prepared and tested for results.

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EXAMPLE 2: HEMOSTATIC PATCH WITH ULTRA-STRENGTHENED MECHANICAL PROPERTIES FOR EFFICIENT ADHESION TO WET SURFACES

Aspects of the following disclosure in Example 2 were also published in the journal Biomaterials, Zheng et al., Biomaterials: Volume 301, October 2023, 122240 (hereinafter “Zheng Biomaterials”).

Bleeding is a potential complication in any surgical procedures, which causes half of the five million traumatic deaths annually (1-3). During major surgeries that target the arteries or highly vascularized organs such as liver, heart or lungs, acute uncontrolled blood loss may lead to increased mortality, morbidity, and rescue times (2, 4). Bleeding wounds are often treated with sutures, wires, staples, or a combination of these interventions (5). However, these conventional wound closure techniques pose additional risks and complications, such as the puncturing of surrounding tissues, which in turn increases the likelihood of infection and further bleeding (5). To overcome this clinical impediment, various types of polymer-based hemostatic patches and bioadhesives have been developed in recent decades; nevertheless, most of them suffer from limited translational success due to the insufficient wet tissue adhesion (6-8), poor mechanical strength (9-11), low elasticity (7, 12, 13), limited biocompatibility (14, 15), and lengthy blood-clotting times (16).

To date, several formulations of bioadhesive hydrogels with hemostatic properties have been developed through the incorporation of hemostatic agents such as laponite (17-19), snake extract (6), tranexamic acid (20), and chitosan (17) within a polymeric network. However, the uncontrolled release of these hemostatic agents into the vascular system may lead to clot formation in distant vessels (21). Alternatively, prefabricated hemostatic patches can be utilized for blood absorption at the wound site, however excess blood absorption within these patches can lead to significant blood loss and morbidity (13). Moreover, the swelling of hemostatic hydrogels (in the presence of blood) can also compromise their adhesion and mechanical integrity (22). In view of this, the prospect of using strong bioadhesives to physically seal the wounds, without triggering the coagulation cascade, is one of the most desirable methods of hemostasis with lower risk of venous thrombosis (13, 23). However, most of the hemostatic bioadhesives developed thus far have demonstrated limited adhesion to wet tissue surfaces (adhesive strength < 40 kPa) (6, 9, 23-36) or they are soft and fragile (6, 7, 23-26, 29, 31, 32, 37-41). An ideal hemostatic bioadhesive should be tough and elastic to impose minimal resistance to the natural deformation of dynamic tissues such as the heart, lungs, and skin. However, the mechanical properties of most developed bioadhesive hemostats have not been well characterized or optimized (42- 44). Since the stiffness of soft tissues ranges from 0.1 to 100 kPa (45), engineering biomimetic bioadhesives with tunable mechanical properties for controlled bleeding in various soft and elastic tissues remains challenging (46). One of the major obstacles in designing these bioadhesives is to achieve a high level of elasticity, strength, stiffness, and toughness simultaneously in a hydrogel network through the incorporation of multiple crosslinking mechanisms that will allow for effective energy dissipation (47-55). To the best of our knowledge, there is no report on engineering a hemostatic bioadhesive with robust adhesion and superb mechanical properties. Furthermore, the bioadhesives are generally susceptible to detachment due to fatigue, which can affect either the material’s cohesive properties or interfacial adhesion forces. The majority of the developed hemostatic bioadhesives lack self-recovery and fatigue resistance characteristics, limiting their applications for clinical use (56).

In addition to improved cohesion and adhesion, engineering ready-to-use hemostatic patches with easy handling, long-time storage, and accessibility is also crucial. Responsive hydrogels hold promise for surgical applications where a trained operator can induce the desired response. However, these smart materials may lack controllability for inaccessible wounds and require additional training for handling, which may reduce their practicality in time-sensitive and life-threatening situations. For instance, majority of these state-of-the-art adhesives require external devices such as light sources (6, 9, 30, 42, 57), co-injectors (7-9, 29, 30, 32, 36, 40), mixing of several components prior to their use (27, 28, 31, 34, 38, 58) or the addition or spraying of a crosslinker (26, 59), rendering them unsuitable for use in emergency settings. Therefore, there is an unmet clinical need for an auto-responsive, ready-to- use adhesive patch that can provide strong wet tissue adhesion, rapid hemostasis, and biomimetic mechanical strength and elasticity for controlled bleeding, sealing, and repairing of internal wounds.

Herein, we aim to introduce a novel elastic, tough, and anti-fatigue hemostatic hydrogel patch that can adhere to wet tissue surfaces within a few seconds of gentle pressing while also supporting the native tissue’s physiological functions. This hemostatic patch, namely APTF hydrogel, was formed based on poly (ethylene glycol) diacrylate (PEGDA) hydrogel interpenetrated with A-hydroxysuccinimide (NHS) conjugated alginate (Alg) and subsequently treated with tannic acid (TA)/ Iron (III) chloride (Fe³⁺) to achieve improved cohesive and adhesive properties. While the covalently crosslinked network of PEGDA formed the primary backbone of the hydrogel, the reversible hydrogen bonding between PEGDA and TA and the electrostatic interactions between Alg-NHS and Fe³⁺ were essential to endow the hydrogel with high elasticity and toughness. In addition, the hydrogel exhibited excellent adhesiveness, achieved synergistically by chemically conjugating NHS to Alg molecules and incorporating mussel-inspired adhesive elements like TA. Adhesion to various substrates including soft tissues and inner organs was characterized in vitro and ex vivo, respectively. The biocompatibility of the engineered multifunctional bioadhesive was assessed both in vitro and in vivo. An in vitro blood clotting assay as well as in vivo rat tail and liver bleeding tests were also conducted to demonstrate the hemostatic properties of our developed APTF patch. In summary, this ready-to-use hemostatic patch exhibited enhanced adhesive and mechanical properties to not only quickly stop bleeding, but also help close and treat internal and external injuries.

Result and discussion

Sequential synthesis of the APTF adhesive hydrogel patch comprised of Alg- NHS, PEGDA and TA/Fe³⁺ is demonstrated in Scheme 1 as shown in Figure 9. First, Alg-NHS and PEGDA were synthesized separately. Then, a solution of Alg- NHS/PEGDA (AP) prepolymer solution in Milli Q water was treated under visible blue light (405 nm) for 4 min in the presence of a lithium phenyl-2,4,6- trimethylbenzoylphosphinate (LAP) photoinitiator to form a covalently crosslinked PEGDA network interpenetrated with Alg-NHS (defined as AP hydrogel throughout the paper). A secondary molecular network was then introduced by the addition of TA molecules which facilitated H-bonds formation between TA and oxygen-rich chemical motifs on the PEGDA backbones. Meanwhile, an ionic crosslinking network was formed inside the hydrogel through the incorporation of Fe³⁺ ions capable of interacting with both TA and Alg-NHS, forming the APTF hydrogel patch as an “all- in-one” crosslinked network. Successful conjugation of NHS to Alg was confirmed with proton nuclear magnetic resonance (1H NMR) (Figure Sla-c in Zheng Biomaterials) and Fourier-transform infrared (FTIR) spectroscopy (Figure Sib in Zheng Biomaterials). The appearance of the peak (~ 2.78 ppm) in 'H NMR indicated the presence of succinimide hydrogens (Figure Sla-b in Zheng Biomaterials, highlighted in blue color). The peak around 2.9 ppm could be due to the impurities from N-(3-dimethylaminopropyl)-N-ethylcarbodiimide hydrochloride (EDC) (Figure S2a in Zheng Biomaterials, highlighted in red), the intensity of which increased with increasing EDC concentrations (Figure Sla in Zheng Biomaterials). Peaks around 2.6 ppm could be related to the impurities from the NHS (Figure S2b in Zheng Biomaterials, highlighted in yellow). Peaks around 3.2 ppm (Figure Sic in Zheng Biomaterials, highlighted in orange color) were presumably due to conjugated EDC to Alg before reacting with NHS (forming unstable O-acylisourea intermediate) after activating the carboxyl group (60). To increase the NHS content in the Alg backbone, Alg: EDC: NHS molar ratio was varied from 1 : 1 : 1 to 1 : 1 : 10 to 1 : 1 :20 for the reaction, respectively. An increase in the degree of NHS conjugation was confirmed by the enhancement of the characteristic NHS peak intensity in the 'H NMR spectrum of Alg-NHS (Figure Sla in Zheng Biomaterials). Additionally, increasing EDC from 1 : 1 : 10 to 1 :2: 10 to 1:4: 10 could increase the degree of NHS conjugation as confirmed by the NHS peak intensity (Figure Sla in Zheng Biomaterials). Alg-NHS synthesized with 1 :4:20 was selected for further experiment due to the higher degree of conjugation compared with other reaction conditions. To confirm successful conjugation, the Alg and NHS mixture (1 :20 molar ratio) was washed and analyzed with 'H NMR and Fourier transform infrared (FTIR), showing the absence of unconjugated NHS peaks (Figure S2c-d in Zheng Biomaterials). Successful conjugation of NHS on Alg was further confirmed with FTIR spectroscopy, which showed characteristic peaks at 1780 cm'¹ (carbonyl stretch of NHS), 1704 cm'¹ (carbonyl stretch of NHS), and 1219 cm'¹ (C-N-C stretch of NHS) (Figure Sid in Zheng Biomaterials). The formation of APTF hydrogel was also assessed with X-ray photoelectron spectroscopy (XPS) studies, which confirmed the presence of N (from NHS) and Fe elements in the ATPF hydrogel (visualized as the blue-shaded areas in Figure Sle in Zheng Biomaterials) but were absent in PEGDA hydrogel. The deconvoluted XPS spectrum of the carbon region showed the presence of 0=C-0 (-288.5 ev, from Alg-NHS), O=C-N (-287 ev, from Alg-NHS) and C-0 (-286.5 ev, from PEGDA) in the hydrogel network. Meanwhile, the presence of TA in the hydrogel was confirmed with pi-pi* interactions in the region of 291.8 ev (Figure Slf in Zheng Biomaterials).

The APTF hydrogel was fabricated in the form of a ready-to-use patch, in which Alg-NHS and PEGDA composed the primary polymeric backbone. NHS was conjugated to Alg to provide instant tissue-material interfacial interactions, while TA was also introduced to enrich the crosslinking network and further tune the mechanical properties by establishing both dynamic H-bonding interactions with PEGDA and ion-induced chelation with Fe³⁺. Additionally, the mussel -inspired adhesive characteristics conferred by TA, along with the presence of NHS functional moieties, provided enhanced tissue adhesion. The covalent and noncovalent interactions among various macromolecules in the designed hydrogel were used to tune its physical properties. The molecular interactions inside the synthesized hydrogel, which were comprised of multiple macromolecular moieties, are demonstrated in Figure la. The covalently crosslinked network of PEGDA was the primary backbone of the hydrogel (Figure lai), which formed hydrogen bonding with multiple OH groups presented on TA to create a TA-crosslinked PEGDA network (Figure laii). On the other hand, the addition of Fe³⁺ ions resulted in chelation with the hydrogel network via the facile formation of mono TA/Fe³⁺ complexes. Fe³⁺ forms bis-chelation with TA when the pH is higher than 1.7 (61). In this work, the observed chelation was a mono-chelation as the pH of the treatment solution containing 40% TA and 3% Fe³⁺ was highly acidic (pH =1.5). Since the chelation between TA and Fe³⁺ proved to be the stronger interactions compared with H-bonding (62), the formation of these TA/Fe³⁺ complexes restricted the H-bond formation between TA and PEGDA, decreasing the crosslinking density of the TA/Fe³⁺- crosslinked PEGDA hydrogel (Figure laiii). Meanwhile, the incorporation of Alg- NHS, which competes with TA to interact with Fe³⁺ ions, induced crosslinking between Alg-NHS and Fe³⁺ ions and recovered the H-bonding interactions between TA and PEGDA. This resulted in the formation of a TA/Fe³⁺ crosslinked Alg- NHS/PEGDA or AP network with significantly increased crosslinking densities (Figure laiv). By controlling the multi-mode molecular interactions, the crosslinking mechanisms in the hydrogel network were modulated, leading to the formation of an ultra-strengthened, multi-component hydrogel network.

As shown in the representative SEM images (Figure lav-viii), the pure PEGDA hydrogel showed the largest pore size relative to the other hydrogels, potentially as a result of lower crosslinking density (Figure lav). Following TA crosslinking, the morphology of the hydrogel changed entirely, demonstrated by a large pore size distribution and inhomogeneous crosslinking (Figure lavi). In contrast, the addition of Fe³⁺ to the hydrogel network led to a more homogeneous TA/Fe³⁺ crosslinked PEGDA hydrogel, with a narrow pore size distribution (Figure lavii). To demonstrate the effects of Fe³⁺ in contributing to homogenous crosslinking throughout the hydrogels, we took images of PEGDA hydrogels treated with TA (40%) and TA (40%)/Fe³⁺ with varied Fe³⁺ concentrations. As shown in Figure S3 in Zheng Biomaterials, the TA-crosslinked PEGDA hydrogel without Fe³⁺ showed inhomogeneous crosslinking, resulting in local swelling of the PEGDA/TA hydrogels. This could be due to the diffusion-governed crosslinking process, as the surface of the PEGDA had a higher crosslinking density than the bulk inner part, accounting for the inhomogeneous crosslinked network and hydrogel deformation. Interestingly, this effect was eliminated by adding Fe³⁺, which could decrease the crosslinking between TA and PEGDA due to its chelation with TA. The final APTF hydrogel showed a more packed structure with fewer pores and higher roughness compared with the other hydrogel compositions (Figure laviii).

Native soft tissues exist within a wide range of mechanical properties (ranging from 3 kPa for the kidney to 100 kPa for the skin) (63). Therefore, engineering bioadhesives with controllable stiffness while maintaining high toughness is critical for sealing various internal organs. The mechanical properties of the APTF hydrogel were evaluated through tensile and compression tests. As shown in Figure lb, increasing Alg-NHS concentrations from 0% to 4% and 8% resulted in an increase in Young’s modulus from 44.96±6.64 kPa to 60.89±6.99 kPa, and 84.06±16.30 kPa, respectively. Similarly, the ultimate strength increased significantly from 676.89±66.19 kPa to 951.05±78.96 kPa by increasing Alg-NHS concentrations from 0% to 8% without decreasing ultimate strain (around 900%) (Figure le). This could be due to increasing Alg-NHS concentrations which directly enriched the ionic crosslinking density via crosslinking with Fe³⁺ and provided recovery of H-bonding interactions between PEGDA and TA within the hydrogel. The net effects of increasing PEGDA concentration from 14% to 30% resulted in a significant increase in Young’s modulus from 42.92±1.79 kPa to 121.94±27.71 kPa (Figure 1c), but a decrease in ultimate strain from 1123.86±101.42% to 696.44±43.30% (Figure If). This could be due to the formation of an enriched, irreversible covalently crosslinked network that could not break and heal, leading to the formation of a brittle hydrogel. Moreover, increasing Alg-NHS concentration from 0 to 8% could increase the toughness of the hydrogel more than 1.5-fold from 2832.77±323.66 kJ/m³ to 4696.62±945.54 kJ/m³ (Figure le) which could not be achieved by increasing the PEGDA concentration (Figure If). Mechanical mismatch between the tissue and bioadhesive can cause tissue damage (64), discomfort (65) and induce inflammation (46, 66). In this work, we demonstrated that the stiffness of APTF patch can be finetuned to match with different soft tissues across their range (1-100 kPa) by varying PEGDA and Alg-NHS concentrations to modulate the crosslinking density, while not compromising the high elasticity and toughness of the hydrogel. Therefore, the engineered APTF patch can serve as a platform to meet the sealing of different tissues through varying polymer concentrations. Here, we selected the APTF hydrogel composition with desirable mechanical properties established as 22% PEGDA, 8% Alg-NHS, crosslinked with 40% TA and 3% Fe³⁺. This representative hydrogel is within the range of Young’s modulus of soft tissues (84 kPa), and achieved a strain of 925%, an ultimate strength of 951 kPa, and a toughness of 4697 kJ/m³, suitable for sealing elastic tissues such as heart. We showed that this composition matched the stiffness of porcine heart tissue (Figure S4 in Zheng Biomaterials) and skin tissues (63) and was used for further characterization. For softer tissues such as the lung and kidney, compositions with lowered stiffness could be used. The mechanical flexibility, instant recovery, and ultra-strengthened toughness of the APTF hydrogel are demonstrated in Figure Id. It is shown that the designed APTF hydrogel was capable of undergoing facile manipulation in various fashions, such as stretching, knotting, unknotting, twisting, or any combination of these, without deformation or any damages due to its superb toughness (Video SI in Zheng Biomaterials). The high elasticity of the engineered APTF hydrogel preserved its integrity during applied dynamic stresses and its self-recovery within 3 sec. Additionally, the hydrogel showed self-healing capability, as two pieces of APTF hydrogels remained adhered to each other after one hour incubation at 37 °C in the wet condition and could withstand the applied stress (Figure S5 in Zheng Biomaterials).

In addition to having sufficient modulus and strength to maintain their structural integrity, self-recovery and fatigue resistance are essential characteristics in designing adhesive materials, which enable them to overcome repeated cyclic loads and maximize their lifetime on dynamic organs. In this regard, the resilience of the APTF hydrogel was investigated through a cyclic loading-unloading compression test. As shown in Figure 2a, an increase in the hysteresis loop was observed at higher strain, which means a higher amount of energy was dissipated. Typically, the area of the hysteresis loop is used to measure the energy dissipated per unit volume, and the area increases with increasing compressive strain (67). At lower strain of 10% and 20%, 0.003±0.003 kJ/m³ (3.3% of the total input work) and 0.049±0.019 kJ/m³ (6.9% of the total input work) were dissipated. At a higher strain of 50% and 60%, 1.313±0.055 kJ/m³ (11.08 % of the total input work) and 2.418±0.095 kJ/m³ (11.23% of the total input work) were dissipated (Figure 2b). The representative cyclic strain and stress curves in Figure 2c demonstrate that the APTF hydrogel could withstand and instantaneously recover from a large strain of 50% even after 100 cycles of loading and unloading. The consistent hysteresis loops over 100 cycles indicated no significant plastic deformation or effective energy dissipation. The compressive modulus and dissipated energy (around 11% of the total input work) showed no significant drop after 100 cycles, demonstrating significant resistance to fatigue (Figure 2d and S6 in Zheng Biomaterials). The energy loss of the APTF hydrogel after 100 cycles was even lower than that of an elastic hydrogel synthesized with an elastin-like polypeptide (30% energy loss after 10 cycles) (68). Overall, the APTF hydrogel demonstrated outstanding elasticity. While significant hysteresis is anticipated for the ions crosslinked Alg hydrogel (69), the synthesized APTF hydrogel exhibited low hysteresis and dissipated energy of -1.35 kJ/m³, presumably due to the joint effects of highly elastic PEGDA hydrogel and recoverable H-bonded hydrogel network (70) between PEGDA and TA, along with lowered crosslinking density between Alg-NHS and Fe³⁺ compared with Alg and Fe³⁺.

In the synthesis of adhesive hydrogel, increasing ultimate strength often leads to a reduction in biomaterial’s stretchability, and improving toughness may lower the stiffness (47-55). The poor elasticity and mechanical strength as well as unmatched stiffness of adhesive hydrogels may limit their clinical applications (e.g., sealing dynamic organs). For example, our group previously engineered a photocrosslinkable, gelatin-based hemostatic bioadhesive with improved adhesiveness compared with gelatin methacryloyl (GelMA) which was tested on porcine lung tissue. However, the Young’s modulus of the material ranged from 172.5 ± 9.2 to 377.9 ± 60 kPa (71), limiting its applications for sealing soft and elastic tissues. In addition, the elasticity and ultimate strength of the resulting hydrogel were not characterized. In another study, Qiao et al. developed a mussel-inspired supramolecular hydrogel composed of chitosan, silk fibroin, and tannic acid with enhanced adhesion and hemostatic properties but the material demonstrated primarily plastic behavior (ductility) due to a lack of covalent crosslinks (12), limiting its utilization in sealing soft and dynamic organs which require high elasticity (10). Further, Lu et al. developed a blue-light- activated cellulose-based adhesive hydrogel for hemostasis, but the material demonstrated insufficient mechanical strength (< 40 kPa), elasticity (<120%), and tunability (9). In addition, Zhu and co-workers prepared a bioadhesive with hemostatic ability based on the co-delivery of polysaccharides and peptide dendrimers, yet it demonstrated weak mechanical properties with an ultimate strength of less than 30kPa and a strain of around 70%. The bioadhesive also faced obstacles in balancing strain and ultimate strength; while increasing the crosslinking density at a higher polymer concentration improved the ultimate strength from 5 kPa to 27.7kPa, it significantly reduced the strain from 65% to 35% (7). Figure 2e shows the mechanical characteristics of previously engineered TA-based hydrogels which are classified across four regions (70, 72-79). The pink region denotes hydrogels with high brittleness featuring low mechanical strength and stretchability, the orange region contains those with high stretchability but low stiffness and ultimate strength. Further, the region in blue shows a midpoint between the two previously mentioned regions comprised of middling stretchability and mediocre toughness. In comparison, due to its molecular design, the synthesized APTF hydrogel (denoted as the red region) did not compromise on stiffness, while maximizing toughness and retaining a high level of stretchability.

Bioadhesives have been studied intensively due to their numerous potential advantages including ease of use, wet-sealing capability, and minimal tissue damage (80). However, challenges exist in conferring a suitable level of biodegradability that persists throughout the wound-healing process and at the same time maintains minimal swelling. The in vitro degradation study in Dulbecco's phosphate buffered saline (DPBS) showed that the engineered APTF hydrogel completely degraded in 2 months (Figure S7 in Zheng Biomaterials). This duration is appropriate for wound healing which generally occurs over 4-6 weeks, while chronic wounds can take even longer (> 3 months) (81, 82). In addition to degradation, the swelling properties of the bioadhesive should be studied and optimized as the pressure changes and compression caused by hydrogel swelling may cause tissue damage by uptaking auxiliary water while also weakening the mechanical properties of the crosslinked networks (83-86). In this study, the APTF hydrogel exhibited minimal swelling (Figure 2f) due to its high crosslinking density and added hydrophobicity resulting from the incorporation of TA (70). In contrast, the AP hydrogel without TA/Fe³⁺ crosslinking demonstrated dramatic and rapid swelling with more than 1000% expansion within 10 hours. In addition to its flexibility, toughness and minimal swelling, the APTF hydrogels could be formed into tubular and complex 3D shapes (Figure 2g) using molding method that can be potentially utilized for a wide range of biomedical applications such as a tubular grasper for vascular anastomosis (87). Furthermore, we explored the possible printability of the developed adhesive hydrogel. Such capability would remarkably extend the applicability of the engineered system for use in more complex and challenging biomedical applications. Benefiting from the proper and adjustable viscosity of AP bioink, it could be extruded easily without bath support and rheology modifiers. Most of the biopolymers for 3D printing need a support bath as they cannot hold their shape without crosslinking or excessive fillers, and once crosslinked, they cannot be extruded for printing (88). However, the extra support structure not only affects the printing efficiency, but the problems caused by the removal pose another concern (89). As shown in Figure 2h, the AP bioink could be printed layer by layer while effectively holding the structure which cannot be achieved by using a molding method. After 3D printing, subsequent treatment of TA/Fe³⁺ could significantly improve the mechanical properties of the printed structure so it could be more stretchable (Figure 2i). The viscosity of AP bioink could be increased by increasing the concentration of Alg-NHS (Figure 2j), which allowed for a more efficient bath- free and additive-free printing procedure as it held its shape completely in situ while could be extruded freely (Figure 2ki). As a proof of concept, we used the AP bioink to print a 1 cm thick tube. After printing the tube, we crosslinked it to further stabilize it into AP hydrogel (Figure 2kii-iii).

The ability to achieve strong and instant adhesion to the wet tissue surfaces is a key limiting factor for the clinical success of bioadhesives, and thus represents the most difficult yet pertinent challenge to overcome. We demonstrated that the APTF hydrogel possessed strong tissue adhesion in which NHS and TA simultaneously generated robust and instant bonding with various wet tissue surfaces, internal organs as well as various substrates (Figure 3). It was found that the incorporation of Alg- NHS in the PEGDA/TA hydrogel significantly increased the adhesion strength (Figure 3a). Notably, adding Fe³⁺ could further enhance adhesion as Fe³⁺ effectively crosslinked with Alg-NHS and prevented Alg-NHS from leaching out from the network during crosslinking. We also demonstrated the positive effect of PEGDA on improving shear strength. Increasing PEGDA concentration from 14% to 22% resulted in a 1.7-fold increase in the shear strength. This is due to the enhanced H- bonding interactions between PEGDA and TA, leading to the incorporation of a higher amount of TA in the network during crosslinking. However, further increasing PEGDA concentration from 22% to 30% did not affect the shear strength (Figure 3b), presumably due to the surface saturation of catechol or pyrogallol moieties of the hydrogel. We further demonstrated that the strong adhesion of the APTF hydrogel could not be achieved by the PEGDA/TA hydrogel in the absence of Alg-NHS even with an increased PEGDA concentration (38%) (Figure S8), emphasizing the synergistic effect of Alg-NHS and TA on improving adhesion strength. To highlight the superior adhesion of the APTF hydrogel, we compared its shear strength with one of the strongest commercial bioadhesives, Histoacryl®, which consists of //-butyl -2- cyanoacrylate. As shown in Figure 3b, the APTF hydrogel exhibited significantly higher shear strength compared to the commercial cyanoacrylate glue. The excellent elasticity and immediate adhesiveness of the engineered APTF bioadhesive to porcine skin are demonstrated in Figure 3c, where the hydrogel remained firmly adhered to the fresh porcine skin under a 358 kPa applied stress after 10 sec of pressing. We further sandwiched the patch between silicone-coated magnetical weights (60 g) and fresh porcine tissue, followed by 1 min pressing time, we immediately lifted the porcine skin with a tweezer and immersed them into DPBS, the hydrogel firmly attached to both the skin and weights, demonstrating its wet adhesive properties (Figure 3d).

To demonstrate the adhesiveness of the APTF hydrogel, we used three American Society for Testing and Materials (ASTM) methods including the lap shear test (ASTM F2255), 180-degree peel test (ASTM F2256) and wound closure test (ASTM F2458) to characterize the engineered bioadhesive. The adhesiveness in each test was determined by pressing down for one min, followed by 15 min of incubation time at 37 °C oven under wet conditions to simulate human body condition. As shown in Figure 3e-f, the APTF hydrogel exhibited an adhesive strength of 77.28±3.03 kPa in the lap shear test and 402.37± 68.76 kPa in the wound closure test, an interfacial toughness of 99.23±7.71 J/m² in the 180-degree peel test (Video S2 in Zheng Biomaterials). These values were significantly higher than the commercial glues such as Coseal, DuraSeal, Histoacryl, BioGlue, and Tisseel reported to be less than 40 kPa in their shear strength (11). Additionally, the adhesive strength of the APTF hydrogel was also higher than the state-of-the-art bioadhesives reported so far (90-93). APTF bioadhesives also exhibited strong adhesion to a wide range of internal wet tissue surfaces including muscle (27.43±11.27 kPa), heart (16.91±6.57 kPa), small intestine (12.61±3.22 kPa), liver (9.78±0.65 kPa), and lung (5.59± 0.84 kPa) (Figure 3g). The differences observed in the adhesive strength of the patch on various tissues are presumably due to the differences in texture, wetness, the density of functional groups presents on the surface, and stiffness of the tested tissues (94). The observed differences are similar to the adhesive strengths reported in the literature for other bioadhesives to the same tested tissues (11, 94, 95). It is worth noting that increasing the incubation time from 15 min to 2 h at 37 °C in wet conditions could significantly increase the adhesive strength from 77.28±3.39 kPa to 143.9±16.74 kPa respectively (Figure 3h) as it allowed for more time to form interactions and enhance the bonding (96, 97). This also confirmed the long-term adhesive properties of the hydrogel, as longer time scales favored the formation of more stable interactions due to in situ auto-oxidation of catechol/pyrogallol, forming highly reactive quinones which could covalently bond to the tissue (98). In fact, most of the engineered ready-to-use adhesive patches reported thus far required significantly longer incubation time for achieving higher adhesive strength (11, 96, 97, 99). For example, Yuk et al. developed a dry double-sided tape for adhesion to wet tissues, which demonstrated a shear strength of 120 kPa, but a 24 h incubation time was allowed prior to the adhesion tests (11).

The APTF hydrogel was developed as a double-sided bioadhesive patch, where one side could readily adhere to tissues while the other could adhere to a surgical device, such as a pacemaker or a titanium implant. Our engineered APTF hydrogel demonstrated instant and high adhesive strength (> 43.7 kPa) to a wide range of substrates including stainless steel, polydimethylsiloxane (PDMS), glass, polycarbonate, acrylic sheets, titanium, and polyamide (Figure 3i). In comparison, a previously reported dry double-sided tape could only adhere to selected substrates that were functionalized with an amine group for improved adhesion (11). Our APTF hydrogel design eliminated the need for surface modification of the substrates as TA, which adopts a mussel-inspired adhesion mechanism, could form a wide range of non-covalent interactions with different surfaces (61, 100). As an adhesive hemostat, it is critical to achieve adhesion to bleeding tissue as the blood can foul the material and therefore drop its adhesion to the tissue surfaces (101). Our APTF hydrogel could adhere to the lung and liver tissues covered with blood with a shear strength of 3.314±0.97 kPa and 2.137±0.45 kPa, respectively, due to its superb wet adhesiveness (Figure S9a-c in Zheng Biomaterials). In contrast, cyanoacrylate completely lost its adhesiveness to bloody tissue with zero shear strength observed (Figure S9b-c in Zheng Biomaterials), proving the efficacy of APTF bioadhesive. As our patch was used in its dry form, it could absorb some blood to expose the adhesive moieties to the tissues (11). We also observed a drop in the adhesive strength of APTF patch in the presence of blood as compared to the patch applied to the tissue in the absence of blood, presumably due to some TA/Fe³⁺ formed interactions with blood instead of tissue (75, 102-105).

To further evaluate the adhesion of our engineered APTF hydrogel, we performed various ex vivo experiments using freshly isolated porcine organs. We showed that APTF hydrogel could effectively seal leakages and hold pressure within the heart (Figure 3ji), intestines (Figure 3jii), aorta (Figure 3jiii), bladder (Figure 3jiv), and stomach (Figure 3jv) within a few sec of pressing (Video S3). The burst pressures for the sealed intestine and heart were 35.72 ±2.74 kPa and 33.37±5.67 kPa, respectively (Figure 3k), which were mainly due to the tissue rupture and not adhesive failure at the interface. The bladder demonstrated a burst pressure of 126.67± 7.26 kPa (Figure 3k), which was due to the combination of tissue rupture and adhesive failure. Aorta tissue showed a burst pressure of 172.10± 24.02 kPa (Figure 3k) with an adhesive failure at the hydrogel -tissue interface. The burst pressure of the engineered APTF hydrogel applied to soft organs was significantly higher than most of the in situ crosslinkable bioadhesives (106-108) and prefabricated bioadhesive patches developed thus far (95-97). In addition, a majority of previously developed adhesive patches experienced cohesive failure due to their limited toughness and mechanical strength. In view of the superb mechanical properties of the APTF hydrogels at different compositions as characterized before, no cohesive failure was observed during all the in vitro and ex vivo adhesive characterizations. Therefore, all the adhesive strength reported here is due to interfacial adhesive failure rather than cohesive failure. The adhesion mechanism of the APTF hydrogel to the wet tissues is shown in Figure 31. Initially, TA interacts with the tissue surfaces through Id- bonding, while NHS, known for its fast reactivity, mainly forms covalent interactions (amide bonds) with the tissue. Subsequent auto-oxidization of TA over time creates quinones to undergo further covalent interactions with the tissue, including Michael addition and Schiff base reactions (63, 77, 109). These covalent interactions collectively account for the durable adhesion of the APTF hydrogel to the tissue, and its long-term retention on tissue surfaces during their physiological functions and activities. Our formulated APTF hydrogel can spontaneously form both noncovalent (due to the incorporation of TA) and covalent interactions (due to the presence of NHS) with tissue surfaces while deforming. To prevent pre-oxidation of TA before applying the patch to the tissue, the APTF patch was formulated in a flexible dry patch form and vacuum sealed in a bag to seclude it from ambient oxygen and moisture (Figure SlOa in Zheng Biomaterials). This cannot be achieved by any catechol/pyrogallol-containing bioadhesive liquid precursors to prevent peroxidation of these adhesive moieties. Therefore, for these types of bioadhesives, hydrogel prepolymers must be prepared fresh prior to use. In our case, we could store APTF hydrogel for long-term use. To assess long-term effective preservation of the patch, after 6-month storage of the vacuum-sealed APTF patch at 4 °C, we performed standard lap shear (ASTM F2255) adhesion tests on fresh porcine skin. As is shown in Figure SlOb in Zheng Biomaterials, the APTF patch, after 6 months of storage, experienced a drop in shear strength from 77.28±3.386 kPa to 44.04±2.045 kPa compared to the freshly prepared APTF patch. But this adhesive strength was still higher than most of the freshly prepared catechol/pyrogallol-containing bioadhesives reported in the literature thus far (9, 26, 33, 93). Additionally, after 6 months of storage, APTF patch showed comparable adhesive strength to cyanoacrylate, one of the strongest commercial bioadhesives. Therefore, it is concluded that APTF patch can be stored for at least 6 months prior to usage.

In vitro hemocompatibility of APTF hydrogel was analyzed by hemolysis of mouse red blood cells (mRBCs). For comparison, the hemolysis ratio (%) of mRBCs exposed to saline was defined as 0. The hemolysis levels of AP hydrogel and APTF hydrogel were 1.061 ± 0.705 % and 3.711±0.041 %, respectively, lower than the ISO 10993-4 (5%) (110), which provided the bioadhesive patch with promising potential to satisfy the clinical requirement. In contrast, mRBCs treated with pure water were completely damaged as expected, with a hemolysis ratio of 100% (Figure Slla-b in Zheng Biomaterials). These data preliminarily suggest that the APTF hydrogels did not change the blood cells’ morphology and thereby had no influence on their normal functions.

In vitro cytotoxicity was assessed using NIH 3T3 cells in Transwell® cell culture inserts to evaluate the biocompatibility of the APTF hydrogel and cyanoacrylate (Histoacryl®). As shown in Figure 4a, APTF hydrogel could effectively support the growth and spreading of 3T3 cells, with no difference compared with the control group where the cells were cultured in the well-plate. This confirmed the in vitro cytocompatibility of the engineered APTF hydrogel. In addition, a high level of cell viability (>95%) over 5 days of culture was observed through quantitative analysis, further demonstrating non-cytotoxicity of the APTF hydrogel. In contrast, cyanoacrylate killed almost all the cells at day 1 with a cell viability of around 2% and cells could not proliferate over 5 days of culture (Figure 4b) Fluorescent staining of F-actin filaments and nuclei, using Alexa Fluor 488 phalloidin and DAPI (4’,6-diamidino-2-phenylindole), confirmed cell spreading on APTF hydrogel over 5 days of culture (Figure 4c). Cell proliferation was further evaluated with PrestoBlue, a resazurin-based cell viability reagent. When added to cells, PrestoBlue reagent is modified by the reducing the environment of healthy cells and turns red (111). This color change can be detected using fluorescence measurements. As shown in Figure 4d-e, the metabolic activity of 3T3 cells increased promptly over 5 days for both the APTF hydrogel and the control, whereas cyanoacrylate showed no metabolic activity over 5 days of culture. These studies together confirmed the in vitro biocompatibility of the APTF hydrogel.

To evaluate the in vivo biosafety of the engineered APTF hydrogel, subcutaneous implantation was performed in the dorsal connective tissue of Wistar rats. AP hydrogels, without TA/Fe³⁺ treatment, were used as control. Histological studies on the explanted hydrogels showed cellular infiltration in the tissue-material interface (marked with yellow arrows in Figures 4f-g). In accordance with hematoxylin and eosin (H&E) (Figure 4f) and Masson's trichrome (MT) (Figure 4g) stained samples, cell infiltration around tissue-hydrogel interface was observed at day 7 and day 28 (Figure S12 in Zheng Biomaterials). Similar results of hydrogel-tissue integration were also obtained for AP hydrogel (Figure S13 in Zheng Biomaterials). It is worth noting that no long-distance cell migration occurred deep into the hydrogel after 28 days post-implantation which could be due to the lack of arginine-glycine- aspartic acid (RGD) moieties in the hydrogel. Both PEGDA and Alg used in the synthesis of APTF are bioinert materials which will not encourage cellular infiltration and growth within the bioadhesive (112, 113). In this work, APTF hydrogel was introduced as a hemostatic tissue sealant. However, to support 3D cell encapsulation and migration for tissue regeneration, the hydrogel system can be functionalized with RGD moieties as Alg-NHS can readily react with RGD to form Alg-RGD (114). Overall, these results confirmed the biocompatible nature of the engineered hydrogel and its ability to undergo integration with host tissues without significant fibrosis.

Upon hemostasis, inflammation is known to be the second stage of wound healing, triggered during the early stages of the healing process through the infiltration of neutrophils and macrophages (115). In light of this, immunofluorescence (IF) staining of macrophages (CD68) was performed to assess the local immune response. Macrophage invasion at the interface between hydrogels and the subcutaneous tissue was observed for both groups (AP and APTF hydrogels) at day 7, presumably due to the host tissue response but almost disappeared completely on day 28 post-surgery (Figure 4h and S12-13 in Zheng Biomaterials), indicating the in vivo biocompatibility of the APTF hydrogel. TA is known to have anti-inflammatory properties and has therefore been adopted as an essential component in bioadhesives assisting wound healing (34, 116). Moreover, PEGDA and Alg have shown in vivo biocompatibility and have been widely used for biomedical applications (117-120). Therefore, both AP and APTF hydrogels - due to their material compositions - satisfy appropriate in vivo biosafety standards.

To the best of our knowledge, there have been no hemostatic adhesives developed thus far that demonstrated simultaneous strong tissue adhesion and excellent in vitro and in vivo biocompatibility. Most of the strong bioadhesives, such as cyanoacrylates, have been associated with cytotoxic effects and inflammation (14, 15). Other known complications of cyanoacrylate-based adhesives include tissue burns due to exothermic polymerization, high stiffness (much greater than that of the skin or other soft tissues), and poor compliance to dynamic deformation (121, 122). Our engineered APTF hydrogel could introduce a higher level of adhesive strength and hemostatic capability compared to commercial bioadhesives, while also ensuring biocompatibility.

The ability to stop bleeding efficiently is critical for injury management and life-saving as the first stage of wound healing is hemostasis. (123) TA has demonstrated high affinity to whole blood components (e.g., red blood cells and platelet-rich plasma) through the generation of a physical adhesive-blood barrier (TA- blood barrier) (75, 102-105). The strong hemostatic effects of TA/Fe³⁺ have been recently demonstrated as a coating material for a hemostatic dressing (124). Similarly, in our study, the molecular design strategy of incorporating TA/Fe³⁺ into the APTF hydrogel provided an effective hemostasis. In vitro blood clotting test was conducted to investigate the hemostatic performance of the APTF hydrogel by exposing the engineered hydrogel to fresh blood and measuring the clot formation over time. As shown in Figure 5a, the clots formed faster on the surface of the APTF hydrogel as compared to the AP hydrogel and the whole blood as a control. Further, hemoglobin absorption was also reduced from 60.57± 2.4% for the AP hydrogel to 11.65 ± 7.76% for the APTF hydrogel within 28 min (Figure 5b). This decreased absorption could be attributed to the high affinity of the incorporated TA/Fe³⁺ to whole blood components, and its ability to facilitate platelet plugs formation for coagulation (125).

The in vivo hemostatic efficacies of AP and APTF hydrogels were also tested in two different animal models: a rat tail bleeding model (Figure 5c-e) and a liver bleeding model (Figure 5f-g). In the first model, 6 cm of each rat’s tail was first transected. The APTF hydrogel, AP hydrogel, and Surgicel® (as a positive control) were then applied to the wound and the blood was collected for 10 min using a filter paper which was then dried and weighed to measure the amount of blood loss (Figure 5c). We used injury without any treatment as a negative control. Although Surgicel® adopted a different hemostatic mechanism compared with the developed bioadhesive, there is no currently available standard of care that adopts the same hemostasis mechanism as our reported work. Surgicel® (oxidized cellulose) is one of the most frequently used hemostatic patches (126) which is also based on polysaccharide; therefore, we chose it as a control in our study. Additionally, commercial hemostatic agents such as Surgicel® and Fibrin glue are widely served as a control in liver bleeding experiments for newly developed bioadhesives (26, 71, 95). As shown in Figure 5d, the mean blood loss was much lower for the APTF hydrogel (29.97 ± 27.09) as compared to the AP hydrogel (124.61 ± 35.62 mg) and Surgicel® (183.6±29.99 mg). No treatment group demonstrated the most significant level of blood loss at 3733.33±873.69 mg (Figure 5d). Representative images of the collected blood on the filter papers are shown for each experimental group in Figure 5e. The hemostatic properties of the AP and APTF hydrogels were mediated by different mechanisms. In the AP hydrogel, this hemostatic effect could be due to rapid absorption of blood to the material, whereas the hemostasis in APTF hydrogel could be due to the combination effects of the incorporated TA/Fe³⁺ and strong adhesive property of the hydrogel.

To assess the in vivo hemostatic efficacy of the engineered bioadhesive to treat hemorrhages of internal organs, we performed an in vivo liver bleeding experiment (Video S4 and S5 in Zheng Biomaterials). In this animal model, four experimental groups were investigated in accordance with the rat tail bleeding model as mentioned before. Immediately after creating a standardized liver wound (2 mm) using a surgical blade (#11), the hemostatic bioadhesives were applied to the bleeding lesion and gently pressed for 1 min (Figure 5f). The blood mass from the incision site until complete coagulation was collected on a filter paper for 10 min and weighed afterward. The mean blood loss for the AP and APTF hydrogels were 4.73±1.00 mg and 2.77±1.74 mg (Figure 5g), respectively, which was significantly lower than the Surgicel® (6.75 ± 0.65 mg). In addition, no treatment group had a blood loss of 24.00±6.56 mg (Figure 5g). Based on these data, it can be concluded that both APTF (Video S4 in Zheng Biomaterials) and AP hydrogels (Video S5 in Zheng Biomaterials) can effectively stop bleeding. Herein, robust adhesion and an immediate reaction to the blood which triggers coagulation are complementary for rapid hemostasis. Despite the fact that hemostatic patches, such as Surgicel®, can stop bleeding, their limited adhesion to the treatment sites may result in their migration to other vital organs, leading to severe health complications (127). On the other hand, injectable or photocrosslinkable adhesive glues can provide suitable adhesion but can be only used to stop bleeding in small wounds with limited hemorrhage (37, 71, 128, 129). Furthermore, the hydrogel precursors can be diluted or fouled by the blood prior to in situ polymerization, which may prevent effective gelation or significantly prolong the hydrogel curing time as well as compromise their cohesion and adhesion. The developed APTF hydrogel showed superior hemostatic efficiency compared to our previously developed bioadhesive hydrogels based on gelatin methacryloyl- catechol (GelMAC) and Fe³⁺’ whose hemostatic effects could not outperform Surgicel® in the same liver bleeding model (71). To demonstrate the efficacy of the APTF patch for controlling bleeding in deep wounds with a large amount of hemorrhage, we performed liver bleeding experiments on rabbits immediately after euthanasia. A 14 mm long deep laceration was made using a surgical blade (#12) to induce hemorrhage. APTF patch was then applied immediately to the bleeding site and followed by pressing for 30 sec. Hemostasis was achieved afterward and was assessed with filter papers. The APTF patch stopped bleeding with minimal blood collected after a gentle pressing, whereas no treatment induced a large amount of blood loss (Figure S14 and Video S6 in Zheng Biomaterials). We also assessed the adhesion of the APTF patch on the rabbit liver after hemostasis, the patch adhered firmly to the wound (Video S6 in Zheng Biomaterials). These preliminary data suggest the efficacy of the hemostatic APTF patch and highlights its potential in treating large hemorrhaging wounds. In our future work, we will systematically characterize the in vivo hemostatic effects of the APTF patch using larger animals (i.e., pigs) with deeper and more complicated wounds to further evaluate its applicability for clinical translation. Due to the degradation of the APTF hydrogel over time (Figure S7 in Zheng Biomaterials), there is no need for removal after its application on the liver tissues. In addition, we do not expect to observe adhesion between tissues in the body due to the presence of the patch, as the adhesion to the tissue requires the application of uniform pressure to the patch over continuous time (at least 5 to 10 sec), and such pressure does not exist between tissues naturally. Additionally, as the hydrogel degrades over time, adhesion would deteriorate over time. Since the adhesion is facilitated through pressing, this may provide the opportunity for the clinicians to do multiple readjustments before applying pressure to eliminate the need for bioadhesive removal due to misplacement. However, this cannot be achieved for cyanoacrylate glue and in situ forming bioadhesive hydrogels (e.g., photocurable adhesives). Additionally, our developed patch has potential for on-demand detachment, as the H-bonds from TA can be cleaved by adding urea (130) at a safe dosage (0.1M), which is known to interrupt hydrogen bonds between TA- containing hydrogels and the tissues for achieving on-demand detachment (70). Urea is a naturally occurring molecule produced during protein metabolism and present in large amounts in human blood and urine (131, 132). Since TA exerts a major role in giving robust underwater adhesion in our hydrogel design, we hypothesize that by adding urea, the adhesion of APTF hydrogel could be significantly weakened for easy removal. Overall, our results confirmed the excellent in vitro and in vivo hemostatic efficacy of the APTF hydrogel driven by the synergy of strong adhesiveness and the incorporation of TA/Fe³⁺ complexes.

We sought to address the currently unmet clinical need in the development of a novel, ultra-strong, anti-fatigue hydrogel capable of sealing bleeding internal wounds and adhering robustly to medical device surfaces. Multiple crosslinking mechanisms and smart molecular design strategies endowed the APTF hydrogel with selective mechanical tunability and high toughness, allowing the material to maintain its integrity while also accommodating dynamic organ function and activity. The resulting hydrogel demonstrated strong adhesive properties due to the synergistic effects achieved by employing different adhesion mechanisms utilizing NHS and TA. This ultimately enabled the APTF hydrogel to adhere to wet tissue within a few sec of gentle pressing without the need for using an applicator or other external devices. In addition, in vitro and in vivo biocompatibility tests ensured its safe utilization in biomedical fields. The robust sealing capabilities of the APTF hydrogel, in addition to its high affinity towards blood, could efficiently terminate excessive bleeding. Thus, the engineered bioadhesive holds great promise to be used in emergency healthcare and can improve tissue-device interfacial adherence for monitoring vital signs and disease diagnosis.

Experimental Section/Methods

Materials

Polyethylene glycol (PEG, 20 kDa), alginic acid sodium salt from brown algae (Alg, 30 -100 kDa), tannic acid (TA), hydroquinone, acryloyl chloride, and lithium phenyl-2,4,6-trimethylbenzoylphosphinate photoinitiator (LAP) were purchased from Sigma- Aldrich. 1 -(3 -Dimethylaminopropyl)-3 -ethylcarbodiimide hydrochloride (EDC) and A-Hydroxysuccinimide (NHS) were obtained from TCI chemicals. Anhydrous iron (III) chloride was purchased from Alfa Aesar. Dulbecco’s modified Eagle medium (DMEM) was purchased from Cellgro (Manassas, VA). Fetal Bovine Serum (FBS) and Dulbecco’s phosphate-buffered saline (DPBS) were obtained from HyClone (Logan, UT). Penicillin/streptomycin (Pen-Strep) and a Live/Dead viability kit were purchased from Invitrogen, Thermo Fisher Scientific.

Synthesis of poly (ethylene glycol) diacrylate (PEGDA)

First, PEGDA was synthesized in toluene via the chemical conjugation of PEG and acryloyl chloride as previously described (133). Briefly, 30 g of PEG (20 kDa) was dissolved in toluene and mixed with 2.5 mL of trimethylamine. Then, 1 mL of acryloyl chloride and 10 mL of dried toluene were added and reacted for 2 hours. Afterward, the reaction mixture was run through a silica bed and filtered into a flask with 200 pL of 30-50 ppm hydroquinone solution in acetone. The produced PEGDA was precipitated in hexane and stored at -80 °C for future use. Proton nuclear magnetic resonance (^JH NMR) was used to verify acrylation, with efficiency calculated at 93%.

Synthesis of Alg-NHS

Alg-NHS was prepared from Alg by esterifying their native carboxyl (- COOH) functional groups. First, 1 gram of Alg was dissolved in Milli-Q water. Then various molar ratios of EDC and NHS were added to the solution. The mixture was allowed to react for 3 h at 45 °C, with the final product being precipitated by the addition of cold ethanol. Afterwards, the product was washed completely with pure ethanol and lyophilized.

APTF hydrogel synthesis First, a 0.5 mg/mL solution of LAP photoinitiator in Milli Q water was prepared. PEGDA and Alg-NHS with varied concentrations were added to the LAP solution and stirred for 1 hour in a dark condition until completely dissolved. After dissolving, the solution was slowly cast in a polydimethylsiloxane (PDMS) mold and introduced to ultraviolet (UV)/visible light (405 nm) for 4 min to polymerize the solution and form the Alg-NHS/PEGDA (AP) hydrogel. Next, solutions of 40% TA and 3% Fe³⁺ (w/v) in Milli Q water were prepared and filtered with a 0.22 pm membrane. The AP hydrogels were then immersed in the TA/Fe³⁺ solution and crosslinked for 24 h at room temperature. Afterward, the hydrogels were extracted from the TA/Fe³⁺ solution, rinsed thoroughly with Milli Q water to remove uncrosslinked TA/Fe³⁺ solution, and vacuum dried overnight. After taking out the patches, the APTF patches were vacuum-sealed and stored at 4 °C for future use. PDMS molds with different shapes (star, tubular, hexagon and UCLA logo) were used to fabricate APTF hydrogels.

Characterizations of Alg-NHS and hydrogel compositions

^JH NMR experiment was used to confirm the chemical functionalization of Alg using a 400 MHz Bruker AV400 spectrometer. The samples were prepared by dissolving 10 mg of the dried polymers in 1 mL of D2O. To confirm successful conjugation, 'H NMR was performed on Alg and NHS mixture at different NHS concentrations (0.5 mg/mL and 1 mg/mL), Alg mixed with NHS (at 1 :20 molar ratio) followed by washing and pure EDC. Fourier transform infrared (FTIR) spectroscopy experiment was operated on a PerkinElmer Paragon 1000 FTIR Spectrometer in the spectral range of 500-3800 cm’¹. The samples were lyophilized in a freeze dryer, and 2-5 mg of Alg, Alg-NHS, Alg and NHS mixture followed by washing were pressed into KBr (200 mg) tablets. The morphology of hydrogels was observed using an SEM (ZEISS Supra 40VP SEM). PEGDA, PEDGA/TA, PEGDA/TA/Fe³⁺, and APTF hydrogels were cross-sectioned, sputter coated with a very thin layer of goldpalladium (Au-Pd) alloy for 30 s, and then examined by SEM. An X-ray photoelectron spectrometer (AXIS Ultra DLD instrument) was used to analyze the chemical composition of the PEGDA and APTF hydrogels. A monochromatic Al Ka X-ray at 15 kV and 10 mA was used as an excitation source. The neutral C is peak (C-C (H) set at 284.6 eV) was used as a reference for charge correction. Rheological investigations were performed by Modular Compact Rheometer MCR302 to characterize the viscosity of AP solutions with varied Alg-NHS concentrations. The results were obtained by linking the measuring system PP08 with diameter of 8 mm to the rheometer. Each measurement (n=3) was carried out by loading a fresh sample in the 1mm gap between the parallel plates and removal of excessive sample. At a given shear rate parameter, ranging from 1 to 100 s ' with 30 measuring points, the relationship of viscosity and shear stress as a function of shear rate was recorded.

In vitro mechanical characterizations of APTF hydrogel

APTF hydrogel samples were equilibrated in DPBS at 37 °C for 1 h to characterize the mechanical properties of the swollen hydrogels. Tensile properties of the APTF hydrogels were measured through pure-shear tensile tests on thin rectangular samples (10 mm in length, 3 mm in width, 0.9 mm in thickness, precise dimensions of each sample were measured using a digital caliper) utilizing a mechanical tester machine (Instron 5943). For all tests (n>5), the tensile speed was set to 50 mm/min. Fresh porcine heart tissues were tested using the same method immediately after harvesting the tissue. Ultimate strength was defined as the stress at the failure point with the strain being read from the same locale. For all samples, the Young's modulus was calculated from the slope of 1/6 - 1/5 of the strain in the stressstrain curve while the toughness was calculated as the area under the stress-strain curve. For fresh porcine heart tissues, Young’s modulus was calculated from the initial linear region of the stress-strain curve. For the successive loading-unloading tests, different strains were tested (10%, 20%, 50% and 60%) with a compression speed of 50 mm/min (n=4). A cyclic compression test was performed on up to 50% of the strain with a constant compression speed of 50 mm/min over 100 cycles (n=4). The cylindrically shaped APTF hydrogel (the precise dimension of each sample measured using a digital caliper) was immersed in DPBS during the test.

In vitro swelling study

The swellability of AP and APTF hydrogels (200 uL AP precusor solution, 9.34 mm diameter, 2.45 mm thickness, cylindrical-shaped hydrogel) was defined as the ratio of weight change to the initial weight of the hydrogel after incubation in DPBS at 37 °C at varied time intervals (n=5). The hydrogels used in the study were created using the protocol listed in the previous section and initially weighed following submersion into DPBS at 37 °C. Excess DPBS was then gently removed using a disposable wipe, and the wet weight of the hydrogels was measured at different time points for up to 72 hours.

In vitro degradation study

APTF hydrogels (100 uL AP precusor solution, 7.18 mm diameter, 1.94 mm thickness, cylindrical-shaped hydrogel) were formed as mentioned before. Samples were incubated in DPBS at 37 °C (n=5). At each time point (0, 10, 26, 36, 45 and 62 days), samples were removed, freeze-dried, and weighed with media being refreshed at each time point. The degradation percentage was calculated based on the weight loss at different time points (Wa) as compared to the initial freeze-dried weight (Wo) following Eq. 1.

Mass residue 100 (Eq.l)

3D Printing of AP bioink

AP solutions were prepared as outlined before and were loaded into a 5 mL syringe affixed to an 18-gauge blunt end needle (BSTEANTM) with an outer diameter of 1.28 mm and inner diameter of 0.84 mm. The syringe was loaded onto an Allevi 3 bioprinter. The pressure was set as 5.0PSI to allow for proper flow rate of the bioinks and printed into different structures with a layer height of 0.5 mm with a custom STL file. The printed structures were then exposed to light (405 nm) to crosslink for up to 1-3 min depending on the shape and the size of the structures. Then, the 3D printed AP hydrogel was crosslinked with TA/Fe³⁺ as described before to form APTF hydrogel.

In vitro adhesive characterizations

For demonstrative purposes, one side of the APTF patch adhered to a piece of wet porcine skin after 10 s of pressing while the other side was clipped with two magnets (60 g in total). For the adhesive lap shear (ASTM F2255), 180-degree peel (ASTM F2256), and wound closure (ASTM F2458) tests, fresh porcine skin tissue surfaces were completely wetted with DPBS before applying the adhesive patch. After adhering the patch to the tissues, it was firmly pressed for 1 min (applied pressure: 2.06 kPa) before being placed in a 37 °C incubation oven for 15 min to equilibrate and simulate human body temperatures. Afterward, a mechanical testing machine (Instron 5943) was used to test the adhered samples (18 mm (length) x 7.5 mm (width)) for the 180-degree peel test and lap shear test, 3.1 mm (width) x 0.8 mm (thickness) for the wound closure test, the precise dimension of each sample was measured using a digital caliper). All tests (n > 5) were conducted with a constant tensile speed of 50 mm/min and the shear and interfacial toughness were determined via dividing the maximum force by the adhesion area and via diving two times the maximum force by the width of the material, respectively. The adhesive strength determined through wound closure was calculated via dividing the maximum force by the cross-sectional area. During the lap shear test and wound closure test, Glass slides layered with cyanoacrylate glue were used to stiffly set the tissues for adhesion to the hydrogels. Lap shear tests were performed on other tissues including fresh porcine muscle, heart, small intestine, liver and lung using the same procedure as previously described. Lap shear tests were performed on different surfaces including stainless steel, PDMS, glass, polycarbonate, acrylic sheets, titanium, and polyamide using the same procedure as described previously. The lap shear test was performed on Histoacryl® composed of n-butyl-2-cyanoacrylate using the same procedure as previously described. To demonstrate the adhesiveness of the APTF patch to the tissue in the presence of blood, lap shear experiments were performed on liver and lung tissues fully covered with blood. The adhesiveness of the APTF patch after 6- month storge was characterized using a lap shear test as outlined before.

Ex vivo adhesive characterizations

Different organs (healthy porcine bladder, small intestine, heart, and aorta) were connected to a water pressure system and Pascal sensor while an incision was introduced to induce leakages. For the heart and aorta, a 5 mm long incision was made using a Miltex Safety Scalpel (#15). For the small intestine and bladder, a 1.5 mm diameter circular incision was made through a needle. For the stomach, a 10 mm diameter circular incision was made by a surgical scalpel. The APTF gel was applied with steady pressure for one minute and left to further adhere for 15 minutes (n = 5). Because these ex vivo tests only required one side of the bioadhesive, the opposite surface was covered with a dampened cellulose sheet to permit sufficient pressure application in the first minute without adhering to the applicator’s glove. Finally, the water system was turned on to introduce a pressure build-up at the site of the patched incisions of the organ. The connected electronic pressure sensor measured the maximum pressure reached before either adhesive or tissue failure.

In vitro hemocompatibility assay

The in vitro hemocompatibility of the APTF hydrogel can be evaluated by hemolysis assay of mRBCs (mouse red blood cells), which were obtained from the blood of the mouse. In detail, 0.3 mL of mRBCs (dispersed in saline) were mixed with 1.2 mL of pure water as a positive control, 1.2 mL of saline as a negative control and 1.2 mL of saline including AP and APTF hydrogel (50 uL AP precursor solution, 5.14 mm diameter, 1.76 mm thickness cylindrical-shaped hydrogel, n = 4). After the incubation (2 h at 37 °C), the treated mRBCs were centrifuged at 3000 rpm for 5 min and the absorbance of the supernatant was measured using a Thermo Scientific NanoDrop One/OneC Spectrophotometer at 541 nm, a wavelength associated with hemoglobin mass and concentration. Hemolysis was computed comply to the Eq. 2. (134)

Hemolysis (Eq.2)

where At is the absorbance of tested samples (AP and APTF hydrogels). A_nc and A_pc are the absorbances of saline and water, respectively.

In vitro cytocompatibility assay

The cytocompatibility of the engineered hydrogels was evaluated through in vitro viability and metabolic activity of NIH 3T3 fibroblasts using Corning® Costar® Transwell® cell culture inserts. Commercial live/dead kits (Invitrogen) and Actin/(4’,6-diamidino-2-phenylindole) DAPI staining (Invitrogen) were used to assess cell viability and proliferation respectively. Metabolic activity was evaluated using PrestoBlue (Life Sciences) assays. NIH 3T3 cells were seeded on the bottom of a 24- wells Transwell permeable (Costar®, 8 pm PET membrane) at 2* 10⁴ cells/cm². AP and APTF hydrogels (10 uL AP precusor solution, 2.83 mm diameter, 0.77 mm thickness, cylindrical-shaped hydrogel) prepared following the previously described protocol and Histoacryl® (10 uL) were placed into Transwell inserts, and 1 mL of growth medium (Dulbecco's Modified Eagle's Medium) was added to each well of the Transwell permeable supports. The well plates were sustained at 37 °C in a humid 5% environment for 5 days with the culture medium being replaced every 48 h.

The viability of 3T3 cells grown on the bottom of well plates was evaluated using a live/dead viability kit according to manufacturer instructions (n = 4). Cells were then briefly stained with 0.5 pL/mL of calcein AM and 2 pL/mL of ethidium homodimer-1 (EthD-1) in DPBS for 20 min at 37 °C. On the first- and fifth-day postseeding, fluorescent imaging was performed using an AxioObserver Z7 inverted microscope. Viable and dead cells were screened by their green and red color, respectively; and quantified using ImageJ software. Cell viability was determined as the number of live cells divided by the total number of cells.

The metabolic activity of the cells was assessed on the first- and fifth-day post-seeding using a PrestoBlue assay (Life Technologies) (n = 6). The 3T3 cells were incubated in 400 pL of 10% (v/v) PrestoBlue reagent in a growth medium for 45 min at 37 °C. Fluorescence was measured using a Synergy HT fluorescence plate reader (BioTek).

F-actin and cell nuclei were used to visualize the spread of 3T3 cells on the bottom of 24-well Transwell permeable supports (n = 4). Cells at days 1 and 5 postseeding were fixed in 4% (v/v) paraformaldehyde (Sigma) for 15 min, permeabilized in 0.1% wt/v Triton X-100 (Sigma) for 5 min, and blocked in 1% w/v bovine serum albumin (BSA, Sigma) for 30 min. Afterward, samples were incubated with Alexa fluor 488 phalloidin for 45 min. Following repeated washes with DPBS, samples were counterstained with 1 pL/mL of DAPI in DPBS for 2 min and fluorescent imaging was completed using an inverted fluorescence microscope (Zeiss Axio Observer Z7).

In vivo biocompatibility assay

The in vivo studies were approved by the IACUC (ARC-2021-113) at the University of California-Los Angeles (UCLA). Male Wistar rats (250-300 g) were purchased from Charles River Laboratories (Boston, MA) and anesthetized by inhalation of isoflurane (~2%). After anesthesia, eight 1 cm incisions were made on the rats’ dorsal skin, and small subcutaneous pockets were made using a blunt scissor. The AP and APTF hydrogels synthesized following the previously described protocol were lyophilized and implanted into the subcutaneous pockets and incisions were closed with 4-0 polypropylene sutures (Ad Surgical) (n = 4). At days 7 and 28 postoperation, the rats were euthanized and the implanted hydrogels were harvested with the tissues surrounding them. Evaluations of inflammatory responses caused by the implanted hydrogels were carried out through histology of the excised hydrogels. After retrieving the hydrogels, they were fixed in 4% paraformaldehyde for 4 h and incubated at 4 °C in 15 and 30% w/v sucrose solution, respectively. The samples were then embedded in Optimal Cutting Temperature (O.C.T) compound, frozen in liquid nitrogen, and sectioned using a Leica CM1950 cryostat machine. The 20-40 pm sections were then mounted on positively charged slides and then were processed for hematoxylin and eosin (H&E) (Electron Microscopy Sciences) and Masson’s trichrome (MT) staining (Sigma) according to manufacturer instructions. Immunofluorescence (IF) staining was also performed on mounted samples as previously reported (135). Anti-CD68 (abl25212) (Abeam) was utilized as a primary antibody, while Goat anti-Rabbit IgG (H+L) antibody conjugated to Alexa Fluor 488 (Invitrogen) was used as a detection reagent and secondary antibody. The samples were then stained using DAPI and fluorescently imaged using a ZEISS Axio Observer Z7 inverted microscope.

In vitro hemostatic analysis

Porcine blood obtained from Yorkshire pig veins was utilized to determine the clotting time of the AP and APTF hydrogels (100 uL AP precusor solution, 7.18 mm diameter, 1.94 mm thickness cylindrical-shaped hydrogel) (n = 4). The engineered hydrogels were first laid into a 48-well plate, then 200 pL of blood was seeded onto each surface of the samples. For each time interval, 10 mL of DPBS was added to dissolve any free red blood cells through gentle shaking and the solution was immediately aspirated and washed. The absorption of the solutions was measured using a UV spectrometer (PerkinElmer, USA) at a wavelength of 542 nm. The hemoglobin content was determined by Eq. 3 (136).

I_s

Hemoglobin absorbance = j- X 100 % (Eq.3)

Is and I_r are the absorbance of the sample and the blank control, respectively.

In vivo rat tail bleeding experiment All animal experiments were approved by the UCLA Animal Research Committee (ARC-2021-113) and conducted in accordance with the relevant guidelines. The animals (Male Wistar rats (250-300 g)) were monitored daily for signs of pain or discomfort during the experimental period. A set of 4 experimental groups were selected for this study which included the no treatment group (injury only) as the negative control, the AP hydrogel, the APTF hydrogel, and the commercial absorbable hemostatic agent, Surgicel® (Ethicon, Cincinnati, OH, USA) to serve as a positive control group (n = 4). Under general anesthesia (1.5% isoflurane in O2), 6 cm of rat tail was transected and hydrogels (AP, APTF hydrogels) and Surgicel® were applied to the wound, followed by 1 min pressing and bleeding was monitored for 10 min. Blood was collected using filter papers and filter papers were dried and weighed to measure the amount of blood loss.

In vivo liver bleeding experiment

In vivo liver bleeding studies were approved by the UCLA Animal Research Committee (ARC-2021-113) and conducted under relevant guidelines. Animals were monitored daily for signs of pain or discomfort during the experimental period. A set of 4 groups were selected for this study as mentioned previously in the rat tail bleeding experiment (n = 4). Under general anesthesia (1.5% isoflurane in O2), a median laparotomy was performed, allowing for a wound retractor and filter paper to be placed appropriately to collect the incision site’s blood loss. A standardized liver wound (1 mm) was made using a #11 surgical blade. Immediately, the sterile samples were applied to the bleeding lesion. The amount of blood absorbed by the filter paper was immediately measured and recorded. Afterward, the abdominal wound was anatomically closed using 5-0 absorbable sutures and 4-0 non-absorbable sutures by closing the peritoneum and the abdominal skin separately.

Rabbit liver bleeding experiments Hemostasis of the APTF patch on the injured liver was evaluated using New Zealand rabbits. Hemostasis was assessed immediately after euthanizing the rabbit to best mimic the in vivo condition. Specifically, a 14 mm incision was made using a surgical scalpel (#12) to induce bleeding. Then, the APTF patch was applied to the wound and gently pressed for 30 sec to enhance the adhesion. Upon removing the pressure, the hemostasis was assessed by collecting the blood with filter papers for up to 5 min. The wound with the same incision without treatment served as a control.

Data analysis Data analysis was carried out using a one-way ANOVA test with GraphPad

Prism 9 software. Error bars represent mean ± standard deviation (SD) of measurements (*p < 0.05, **p < 0.01, ***p < 0.001, and ****p < 0.0001). All the experiments with statistical analysis were repeated for at least 4 times.

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PUBLICATIONS

All publications mentioned herein (e.g. those listed numerically herein) are incorporated herein by reference to disclose and describe the methods and/or materials in connection with which the publications are cited. Publications cited herein are cited for their disclosure prior to the filing date of the present application. Nothing here is to be construed as an admission that the inventors are not entitled to antedate the publications by virtue of an earlier priority date or prior date of invention. Further, the actual publication dates may be different from those shown and require independent verification. The following references include descriptions of methods and materials in this field of technology.

CONCLUSION

This concludes the description of the illustrative embodiments of the present invention. The foregoing description of one or more embodiments of the invention has been presented for the purposes of illustration and description. It is not intended to be exhaustive or to limit the invention to the precise form disclosed. Many modifications and variations are possible in light of the above teaching.

Claims

CLAIMS:

1. A hydrogel composition comprising:

N-hydroxysuccinimide ester (NHS) conjugated alginate (Alg-NHS); poly (ethylene glycol) diacrylate (PEGDA); tannic acid (TA); and Fe³⁺ ions.

2. The hydrogel composition of claim 1, further comprising a photoinitiator agent.

3. The hydrogel composition of claim 1, further comprising a therapeutic agent or an imaging agent.

4. The hydrogel composition of claim 1, wherein the hydrogel exhibits at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa.

5. A method of making a hydrogel composition comprising: combining an alginate conjugated to N-hydroxysuccinimide ester moieties with a poly (ethylene glycol) diacrylate (PEGDA); including in the combination tannic acid (TA); and Fe³⁺ ions; such that the hydrogel composition is made.

6. The method of claim 5, further comprising including in the combination a therapeutic agent or an imaging agent.

7. The method of claim 5, wherein the hydrogel is formed to exhibit at least one material property selected from: an elasticity > 500%, 600%, 700%, 800% or 900%; a toughness >2,500, 3,500 or 4500 kJ/m3; and/or tissue adhesion in a wound closure test of > 200, 300 or 400 kPa.

8. The method of claim 5, further comprising a photoinitiator agent.

9. The method of claim 8, wherein the method comprises exposing the composition to visible blue light so as to activate the photoinitiator agent and form a covalently interconnected PEGDA network interpenetrated with Alg-NHS.

10. A method of adhering a first tissue region to a second tissue region, the method comprising: disposing a hydrogel composition of claim 1 at a site where the hydrogel composition is in contact with the first tissue region and the second tissue region; and forming a covalently interconnected polymer network form the hydrogel composition; such that the first tissue region is adhered to the second tissue region.

11. The method of claim 10, wherein the first tissue region is adhered to the second tissue region in vivo.

12. The method of claim 10, wherein the hydrogel composition is prefabricated as an adhesive patch.

13. The method of claim 10, wherein the hydrogel composition is 3D printed.

14. The method of claim 10, wherein the hydrogel composition inhibits bleeding from the first tissue region as the first tissue region is adhered to the second tissue region.

15. The method of claim 10, wherein the hydrogel composition further comprises a mammalian cell, a therapeutic agent or an imaging agent.