Instant tough adhesion of polymer networks

Significance

This work presents a simple strategy to generate instant tough adhesion between hydrogels using chitosan films. Strong film mechanics and adhesive polymer films with pKa around physiological pH are key for achieving adhesion through interfacial topological entanglement and other non-covalent interactions such as H-bonding and Van der Waals forces. These findings and the simplicity of this technology support clinical applications and have important implications for designing composite hydrogels and interfacing devices with the human body where fast and robust adhesion between gels and other polymeric materials is required.

Abstract

Generating strong rapid adhesion between hydrogels has the potential to advance the capabilities of modern medicine and surgery. Current hydrogel adhesion technologies rely primarily on liquid-based diffusion mechanisms and the formation of covalent bonds, requiring prolonged time to generate adhesion. Here, we present a simple and versatile strategy using dry chitosan polymer films to generate instant adhesion between hydrogel–hydrogel and hydrogel–elastomer surfaces. Using this approach we can achieve extremely high adhesive energies (>3,000 J/m²), which are governed by pH change and non-covalent interactions including H-bonding, Van der Waals forces, and bridging polymer entanglement. Potential examples of biomedical applications are presented, including local tissue cooling, vascular sealing, prevention of surgical adhesions, and prevention of hydrogel dehydration. We expect these findings and the simplicity of this approach to have broad implications for adhesion strategies and hydrogel design.

Hydrogels have found wide utility in the biomedical sciences, with applications ranging from tissue repair to bridging materials for the integration of bio-electronic devices (1, 2). Despite these advances, the ability to rapidly create strong interfacial toughness between hydrogel polymer networks has largely been unresolved (3–7), traditionally relying on the formation of covalent bonds between the adherends (1, 8), which is often accompanied by comparatively weaker interfacial adhesion, complicated application procedures, or longer than desired adhesion times. Recently, a hydrogel adhesion strategy using a liquid chitosan bridging layer (tough adhesive, TA) was reported that does not rely on the formation of covalent bonds (9). This liquid chitosan-based approach exhibits time- and pH-dependent adhesion mechanisms. Over time, chitosan (pKa ~ 6.5) diffuses into the hydrogel surface (pH > 6.5) and becomes deprotonated. Chitosan deprotonation results in hydrogen bonding between polymer chains, leading to their entanglement and formation of molecular interlocks between the bridging polymer and the adherend matrix, together generating a threefold to fourfold improvement in adhesion strength compared to alternative strategies that rely on covalent bond formation alone (9).

A central limitation, however, of liquid-based bridging polymers is that the strength of adhesion is time-dependent, further challenged by the presence of a hydration layer that exists on wet materials and tissues, and typically requires minutes to hours to reach equilibrium (3, 9–11) due to the low diffusivity of macromolecules (12). In contrast to “liquid-based” adhesion strategies, “dry” adhesion relies on rapid absorption of fluids at the hydrogel–hydrogel interface and simultaneous chemical bonding, enabling instant adhesion (13, 14). To generate strong, stable, and robust adhesion, previous dry adhesion strategies required the formation of covalent bonds directly with the adherend, which relied on specific reactive functional groups and/or surface modification of the substrates (13). These requirements typically involve complex pre-treatment of the adherends, the introduction of highly reactive functional groups, or light irradiation which can make clinical applications more challenging (15–18).

To overcome these difficulties, we recently demonstrated rapid hydrogel adhesion to wet tissues by coupling the hydration of a dry tough polymer matrix with liquid chitosan bridging polymers (11). While offering great potential in creating hydrogel-to-tissue adhesion, these liquid-based bridging polymer solutions have been unsuccessful at generating fast noncovalent adhesion between hydrogels or hydrogel-elastomer hybrids (19, 20). Here, we address this challenge by using dry bridging polymer films which generate instant tough adhesion between hydrogel–hydrogel and hydrogel–elastomer surfaces via physical non-covalent interactions including chain entanglements and hydrogen bonding. Our approach avoids the need for complex adhesion procedures and toxic components, demonstrating a simple and versatile method for achieving adhesion between soft polymeric materials.

Results and Discussion

In contrast to previously described liquid-based bridging polymers, that require diffusion into the adherends and subsequent entanglement to generate adhesion (9), we present film-based dry bridging polymers to generate rapid adhesion (SI Appendix, Fig. S1). Specifically, dry chitosan films are used to create adhesion through localized and dense surface entanglement and H-bonding to tough alginate-polyacrylamide hydrogels, a double network hydrogel with extraordinary material toughness (21) (Fig. 1A). Polarized light imaging (SI Appendix, Fig. S2) and elemental mapping via energy dispersive spectroscopy (EDS) (Fig. 1B and SI Appendix, Fig. S3) revealed a homogeneous chitosan film (CF) region at the interface between the adherent tough gels (TG) following application. Instant (<1 s) tough adhesion was achieved between hydrogels under standard peel test conditions (Fig. 1C and SI Appendix, Fig. S4 and Movie S1) and when pulled in tension (Fig. 1D). Unlike previous studies evaluating the performance of other bioadhesives, we observe very sharp periodic peaks in the adhesion curves (SI Appendix, Fig. S4), which could likely be due to the strain-stiffening response of the TG prior to its cohesive failure.

Fig. 1.

Chitosan films generate instant strong adhesion between hydrogels. (A) Schematic highlighting differences between liquid and film-based hydrogel adhesion strategies. (B) EDS-based elemental mapping of the interfaces using liquid chitosan (C) and chitosan films (CFs) applied between two adjacent alginate-polyacrylamide tough gels (TG). In both cases, the bridging layer was rich in chlorine⁻, a surrogate measure for chitosan (chitosan-HCl); (Scale bar: 100 µm.) (C) Applying CF between two TGs results in instant adhesion. (D) Robust adhesion between two gels is demonstrated by their maintained apposition despite high tension. Yellow lines denote the borders of the CF after attachment to the TG.

To evaluate the versatility of this dry film–based adhesion approach, several polymer films based on bridging polymers with different functional groups like amine (-NH₂) and carboxylic acid (-COOH) groups, including poly(amino styrene) (PAS), poly(acrylic acid) (PAAc), and carboxymethyl chitosan (CMC) were also investigated for their impact on hydrogel–hydrogel bonding. PAS (pKa ~ 4.5), a polymer with a high density of primary amines achieved weaker adhesion compared to chitosan, demonstrating that the simple presence of amine functional groups was not solely responsible for the observed strong adhesion with CFs. Attempts to further clarify the role of amine functional groups in the hydrogel adhesion process were unsuccessful, since other amine-containing polymers such as poly(allylamine) (PAA) and polyethyleneimine (PEI) were incapable of generating films after solvent evaporation, which may have been due to their C–C flexible backbones, small side groups, or shorter polymer chain lengths. We next evaluated the performance of polymer films bearing carboxylic acid groups capable of H-bonding including PAAc (pKa ~ 4.5) and CMC (pKa ~ 2 to 4) (22). In contrast to the strong adhesion of the CFs, films generated from PAAc or CMC did not demonstrate any adhesion properties in our hydrogel–hydrogel bonding assay.

In summary, and in contrast to all the other hydrogen-bonding-capable polymeric films described above, chitosan-based films generated instant (<1 s) tough adhesion between adjacent hydrogels and achieved the highest adhesion energy (>3× that of other investigated films) (Fig. 2A). When evaluating the results obtained from the different polymer films employed in our hydrogel adhesion survey, it is important to note that carboxylic acids become charged and lose part of their ability to form H-bonds when the pH > pKa, while amino groups lose their charge and gain the ability to form H-bonds when the pH > pKa. Since all of our hydrogels were made in Hanks’ Balanced Salt Solution (HBSS) buffer (pH ~ 7.4) and only chitosan (pKa ~ 6.5) was able to generate strong hydrogel–hydrogel adhesion under these conditions, these observations suggest that the bridging polymer used as dry films must be able to both form H-bonds and exhibit strong chain–chain interactions within the target pH range of the hydrogel system under consideration.

Fig. 2.

Chitosan films generate tough and stable adhesion. (A) The effect of different bridging polymer films on the adhesion energy between Alg-PAAm gels. Data shown as mean ± SD (n = 4 gels/group), as evaluated by a one-way ANOVA with post hoc t tests with Bonferroni corrections. (B) Application of CFs to TGs led to wrinkling immediately after attachment macroscopically (Top) and under confocal imaging (Bottom). Red indicates TG and green indicates CF. (C) The effect of incubation time in DMEM at 37 °C on the adhesion energy between Alg-PAAm gels. Data shown as mean ± SD (n = 4 to 7 gels/group), as evaluated by a one-way ANOVA with post hoc t tests with Bonferroni corrections. (D) Schematic of instant elastomer attachment (VHB, 3 M) with CFs. (E) Strong adhesion (up to 4,000 J/m²) observed between the TG-CF and the VHB elastomer. (F) Stability of TG attached to VHB elastomer with a CF after 0 and 24 h of incubation in DMEM at 37 °C. Dashed lines indicate comparison to the benzophenone strategy for elastomer attachment as prepared and for fully swollen TGs (15).

CFs exhibited robust mechanical properties in their dry state (SI Appendix, Fig. S5 A–C) which can contribute to their strong adhesion to TGs. The adhesion energy generated between TGs was not affected by chitosan dry film thickness (SI Appendix, Fig. S6), despite differences in wrinkling patterns of the films after placement onto hydrogel surfaces (Fig. 2B and SI Appendix, Fig. S6). With thicker films, wrinkling patterns became minimal, likely due to the inherent bending rigidity of the film, while still achieving strong adhesion (SI Appendix, Fig. S6C). Thicker CFs showed elevated moduli, which could be due to changes in packing during dehydration and different hydration levels from ambient humidity (SI Appendix, Fig. S5B). Because the films are not chemically crosslinked, many of the tested bridging polymer films dissolve once immersed in fluid of neutral pH. However, serial confocal imaging revealed stable CF thickness after application between gels through 100 min, suggesting that the CFs remain as a distinct layer between gels and do not undergo significant swelling (SI Appendix, Fig. S7), despite lacking any extensive inter-chain covalent crosslinking. The TG adhesion to the CFs was so strong in fact that after peel testing, residual TG material was observed on the bridging chitosan surface due to cohesive failure with the TG matrix (SI Appendix, Fig. S8). Following incubation in Dulbecco’s Modified Eagle Medium (DMEM) at 37 °C for 12 h, 24 h, and 48 h, high adhesion energy was maintained, although there was some decline over time due to cohesive failure, highlighting this system’s sustained performance even under challenging conditions (Fig. 2C).

While our initial studies focused on hydrogel–hydrogel adhesion, generating adhesion between hydrogels and elastomers has also been a challenge as traditional methods rely on toxic glues, surface functionalization, or complex procedures (15, 16). Because of the potential impacts in the fields of bioelectronics and wound dressings, we decided to test whether our CFs would be able to couple TGs to elastomers. CFs enabled rapid, robust attachment between TGs and acrylic elastomers (VHB 4905, 3M) potentially through cooperative non-covalent bonding between the dense film and the two adherends (Fig. 2D) (23). Attachment of the TG-CF to VHB produced instant strong adhesion (up to >4,000 J/m²) (Fig. 2E and Movie S2), without the need for chemical treatments, use of ultra-violet (UV) light for crosslinking that may require an excess of 30 min for bonding to occur (15), or use of toxic cyanoacrylates (16). This fast adhesion is likely due to the cooperative contribution of both multiple Van der Waals and hydrogen bonding interactions between the chitosan polymer chains and the polyacrylate network of the VHB. X-ray photoelectron spectroscopy (XPS) survey scans of the VHB surfaces, however, did not show any nitrogen in these elastomeric networks (which is present in both the TG and the chitosan), suggesting that the presence of nitrogen-containing functional groups is not required for CF-mediated adhesion (SI Appendix, Figs. S9–11). Additional characterization studies employing Raman spectroscopy identified the presence of C=O carbonyl groups in VHB (SI Appendix, Fig. S12), which could correspond to either ester or carboxylic acid functionalities, both of which can contribute to some of the non-covalent interactions highlighted here (24, 25). Adhesion to VHB was also stable and remained unchanged following incubation in culture medium for 24 h (Fig. 2F). In addition to its high adhesion strength, we further observed that the simple single-step addition of this VHB elastomeric coating to the hydrogel surface delayed dehydration of the hydrogel (SI Appendix, Fig. S13), suggesting its utility in biomedical application where hydrogels could be exposed to the external environment.

To further investigate the mechanism of CF adhesion to TGs, the potential contributions of electrostatic interactions were first considered. Although alginate carries a negative charge and chitosan carries a positive charge at neutral pH (Fig. 3A), interfacial toughness comparable to that of liquid chitosan after 24 h (9) was achieved instantly in net neutral PAAm-only gels using CF (Fig. 3B and SI Appendix, Fig. S14). This finding highlights that adhesion energies are dependent on gel mechanics and suggests that this strategy could be extended to other hydrogel compositions.

Fig. 3.

pH and entanglement affect adhesion. (A and B) The effect of electrostatics on adhesion. Data shown as mean ± SD (n = 3 to 4 gels/group), as evaluated by a Student’s t test. (C and D) The effect of TG pH on adhesion. Data shown as mean ± SD (n = 3 to 4 gels/group), as evaluated by a Student’s t test. (E and F) The effect of prior CF deprotonation on adhesion. Data shown as mean ± SD (n = 2 to 6 gels/group), as evaluated by a Student’s t test. (G and H) The effect of dangling chain ends and surface entanglement on Alg-PAAm gel adhesion. Data shown as mean ± SD (n = 3 to 6 gels/group), as evaluated by a Student’s t test. Ring structures are used to illustrate the concept of molecular interlock/topological entanglement of polymer chains.

CF-based adhesion assay screening with different hydrogel types, however, revealed that hydrogel-hydrogel adhesion was only observed with other acrylamide-based formulations, including Alg-NIPAm thermoresponsive gels (4) and PAAm-only hydrogels (SI Appendix, Fig. S15) (26, 27). Furthermore, the extent of swelling of TGs also influenced the adhesion strength, suggesting that the density of polymer chains in the adherend, hydration level, and the hydrogel structure/topology could also be playing important roles in the adhesion process (SI Appendix, Fig. S16).

Previous studies have shown that hydrogels with pH below 6.5 exhibited reduced interfacial toughness when using liquid chitosan as a bridging polymer, likely resulting from decreased deprotonation of amine functional groups in the chitosan polymer chains (9, 11). This pH adhesion dependency represents a practical challenge in a biomedical context since human tissues and their adjacent environments can vary widely in pH (28). We therefore tested whether these pH effects were also present in the dry CF-based adhesion strategy. Decreasing the pH of the Alg-PAAm TGs decreased adhesion energy (Fig. 3 C and D), but not TG mechanics (SI Appendix, Fig. S17), suggesting that the pH of the adherends influences adhesion when utilizing the dry CFs. Further investigations revealed that deprotonating the amine groups in the CF through immersion in 1 M NaOH inhibited subsequent hydrogel-hydrogel adhesion (Fig. 3 E and F). Based on these observations, we propose that deprotonation of the CFs before application creates stronger interactions between the chitosan chains, preventing them from interacting with the hydrogel substrates. In addition to pH and electrostatics, the importance of surface entanglement to generate adhesion was evaluated by using gels with different surface topologies (Fig. 3G). Hydrogels cast on hydrophobic oxygen permeable polycarbonate (26, 29, 30), which increases the presence of dangling PAAm chains in the double network structure by interfering with the free radical polymerization process at the surface (31), demonstrated significantly lower adhesion (Fig. 3H). TGs casted on polycarbonate substrates showed some decrease in mechanical properties (SI Appendix, Fig. S18); however, this approach completely abrogated adhesion (Fig. 3).

Since dry CFs enabled unprecedented instant adhesion between Alg-PAAm and PAAm-only hydrogels compared to previous studies (3–7, 9, 11, 20), several proof-of-concept examples for their applications as fast self-adhering biomaterials were explored when integrated with a TG or TA (Fig. 4A). For example, the TG can easily be wrapped around cylindrical objects such as a finger (Fig. 4B), demonstrating their favorable handling characteristics that would be practical in surgical or other clinical scenarios. Due to the high thermal conductivity of water in the TGs, we also explored their capacity to provide local tissue cooling. Following TG application to the surface of a human palm, the temperature of the skin decreased unlike conventional standard-of-care hydrogel (e.g., Tender Care®) (Fig. 4C), likely due to elevated water content in the TG. This property could have clinical implications for patients suffering burn injuries, by maintaining regional cooling, while preventing rapid fluid loss through the wound surface.

Fig. 4.

Diverse applications of self-adhering tough gels. (A) The CFs function both to instantly self-adhere with (1) TGs or (2) TAs for several indications including: topical (skin) and internal (bowel, tendon, nerve, vessels). TG: Alg-PAAm only. TA: Alg-PAAm + liquid chitosan. Topical applications: (B) Application of the self-adhering TG around a finger. (C) The TG resulted in local skin cooling in contrast to TenderCare®. Data shown as individual points before and after application (n = 7 hands/group), as evaluated by a two-way ANOVA for time and treatment. Internal applications: (D–F) The TG + CF is easily wrapped around bowel, tendon, and peripheral nervous tissue to provide self-adhesion but anti-adhesive properties to underlying tissue and surrounding organs. (G and H) High-frequency ultrasound imaging confirmed apposition of the gels to tendon and nerve tissue. (Scale bar: 1 mm.) (I) Application of the CF over a TG used as an aortic sealant to increase its strength (C is liquid chitosan). (J) The CF is easily applied over the surface of the TG after adhesion to porcine aorta and withstood cyclic loading. (K) The effect of CF thickness on maximum burst pressures. Data shown as mean ± SD (n = 3 samples/group), as evaluated by a one-way ANOVA with post hoc t tests with Bonferroni corrections.

The formation of fibrotic adhesions following surgery or injury is common and can have devastating consequences, and as such, their prevention represents another unmet clinical need. While commercial technologies such as Seprafilm® (Baxter, Deerfield IL) provide a hydrophilic surgical adhesion barrier, their widespread use has been limited by their poor mechanical properties. Furthermore, these materials can rapidly degrade in aqueous media, and as such, are contraindicated in many operations where postoperative adhesions are a primary concern, such as in bowel anastomosis cases. Since TGs alone do not adhere to tissues, our self-adhesive TG-CF system could be useful for internal applications that require gel-to-gel adhesion but non-adherence to the underlying tissues, therefore maintaining tissue planes while also promoting physiologic gliding (i.e., anti-adhesion) with adjacent structures. To explore the utility of this approach, we have successfully demonstrated that our self-adhesive hydrogel construct can be easily wrapped around organs such as the bowel (Fig. 4D), tendons (Fig. 4 E and G), and peripheral nerves (Fig. 4 F and H) highlighting their potential application as anti-adhesion barriers with surrounding tissues (32).

In an extension of our previous applications described above, we further explored the utility of our TG-CF system in providing mechanical reinforcement for hydrogels employed as internal sealants for fluid leaks in the body. In this ex vivo study and using liquid chitosan to facilitate TG-tissue adhesion, we investigated whether the addition of a CF could serve as a clinically relevant and biocompatible hydrogel backing material to increase peak burst pressures (Fig. 4I). The CF was applied to the liquid chitosan-adhered TG on the surface of a swine aortic arteriotomy model and withstood thousands of simulated cardiac cycles, highlighting a simple strategy to increase its burst pressure resilience (Fig. 4J). Peak burst pressures were approximately twofold higher with the addition of the CF and higher than peak values reported from using other hemostatic bio-adhesives (33). Furthermore, the measured burst pressure could be further increased by employing thicker (higher stiffness) CFs (Fig. 4K).

In this study, we have introduced a simple and versatile strategy for generating strong, robust, tough adhesion between polymer networks, including tough hydrogels and acrylic elastomers. By relying on non-covalent interactions mediated by dry chitosan bridging polymer films, our system does not require any surface pre-functionalization or reactive functional groups for covalent bond formation. We have demonstrated and envision several potential biomedical applications that could be enabled by this adhesion strategy including the on-demand tunability of hydrogel moduli in the surgical theatre, the creation of self-adhesive tissue wraps, and the rapid encapsulation of flexible electronics and sensors for medical diagnostics. Inspired by these results, we expect these findings and the simplicity of our approach to have important implications for designing composite hydrogels, and interfacing devices with the human body where fast and robust adhesion between hydrogels and other polymeric materials is required.

Methods

Tough Hydrogel Synthesis.

Alginate-PAAm tough hydrogels were synthesized by mixing one syringe containing a 14 mL solution of 2.2% sodium alginate (FMC BioPolymer LF10/60) and 13.5% acrylamide (Sigma, A8887) in HBSS (Gibco), 50.4 µL of 2% N,N′-methylenebis(acrylamide) (Sigma, M7279), and 11.2 µL of TEMED (Sigma, T7024), with a second syringe containing 316.4 µL of 6.6% ammonium persulfate (Sigma, A9164), and 267.4 µL of 0.75 M calcium sulfate dihydrate (Sigma, 31221) (34). The gel was cast into glass molds (110 × 15 × 1.7 mm³) sealed on top and bottom with glass and left to crosslink for 24 h. For gels with dangling chains on the surface, polycarbonate molds (110 × 15 × 1.7 mm³) were used. After 24 h, tough gel strips were removed from molds and stored in sealed plastic bags at 4 °C. For hydrogels with lower pH, the TG pH was changed through swelling gels in a buffer of different pH supplemented with calcium, followed by dehydration to baseline. The dehydration/wash process was completed twice to ensure proper buffer exchange.

Polyacrylamide Hydrogel Synthesis.

A similar procedure was used as in the tough hydrogel synthesis, but in this case, no alginate was added to the 13.5% acrylamide stock solution and no calcium was added to the second syringe. The gels were cast into glass molds (110 × 15 × 1.7 mm³) sealed on top and bottom with glass and left to crosslink for 24 h. After 24 h, the tough gel strips were removed from molds and stored in sealed plastic bags at 4 °C.

Synthesis of Other Common Hydrogels.

Other common hydrogels were synthesized in glass molds (110 × 15 × 1.7 mm³) following standard procedures reported in the literature including Alg-NIPAM (4), PVA (16), and PHEMA (16).

Generation of Polymeric Films.

Polymer films (chitosan, PAS, pAAc, CMC) were generated by first dissolving polymers in nanopure water at 2% and 4% concentration and casting 10 mL of solution into molds on glass (6.5 × 9.5 cm). As-received Ultrapure chitosan HCl (HMC, #54046) solution was placed in a 55 °C oven for 6 h to dry and then stored in sealed bags with desiccant for future use. N,O-carboxymethyl chitosan (HMC, #44002), pAAC (Sigma, 323667), and PAS (Polysciences, #02823-1) films were prepared by letting the polymer solution evaporate at room temperature until dry films were obtained. NaOH-treated CFs were prepared by submerging the dry chitosan film in a 1 M NaOH solution for ~1 h. The films were then removed from the solution, rinsed with nanopure water, and left to dry under a vacuum.

Chitosan Films Mechanical Testing.

Pure shear tests were carried out to measure the matrix toughness. In brief, rectangular specimens (20 × 5 mm²) were tested in tension (Instron 3342, 50 N load cell) at a loading rate of 100 mm/min. From the stress–stretch curves, the matrix maximum stretch, maximum stress, and fracture toughness were calculated (21) using custom MATLAB code.

Adhesion Testing.

Adhesion energy (also known as interfacial toughness) was measured with 180° peeling tests (Instron 3342) under uniaxial tension (100 mm/min). A chitosan film (60 × 15 mm²) was placed between two hydrogels and compression (5.5% strain) was briefly applied for <30 s. The back of TG was also bonded to a rigid polyethylene terephthalate (PET) film with cyanoacrylate (Krazy Glue), to limit deformation to the crack tip, and thus all the work done by the machine would be equal to the energy dissipated at the crack tip. The free ends of TA and the substrate were attached to acrylic sheets, to which the machine grips were attached. A mechanical testing system (Instron 3342, 50 N load cell) was used to apply unidirectional tension, while recording the force and the extension. The loading rate was kept constant at 100 mm/min. The adhesion energy (also known as interfacial toughness) was measured as two times the peak value of the ratio of the force and width (3).

Confocal Microscopy.

The adhesive interface and chitosan diffusion profiles for both the hydrated and dehydrated CFs were studied by using confocal imaging and fluorescently labeled chitosan (FITC Chitosan). FITC Chitosan was synthesized by reacting fluorescein isothiocyanate (Sigma, 1245460250) with chitosan. Briefly, 1 g of ultrapure chitosan HCl was dissolved in 100 mL of 0.1 M acetic acid and 100 mg of FITC was dissolved in 100 mL of anhydrous methanol at 1.0 mg/mL in separate flasks. The FITC solution was then slowly added to the chitosan solution under continuous stirring. After ~3 h, the reaction was quenched by slowly adding NaOH (0.5 M) to increase the pH to ~10, precipitating the fluorescently labeled chitosan. The solutions were then centrifuged, and dIH2O was added to the precipitate after discarding the supernatant. These purification steps were repeated several times until FITC was no longer observed in the supernatant. The final product was dialyzed against acidic water for ~2 d and freeze-dried. The FITC Chitosan was transferred to a container with aluminum foil to protect it from light and stored at 4 °C until further use.

To evaluate the gel–film interface and morphology, the adhesion procedure described previously was repeated using a 1:5 mixture of FITC:no-FITC chitosan as the bridging polymer for the film and fluorescently labeled gels with 0.1% nile blue acrylamide (#25395-100, Polysciences). Whole samples were used, or sagittal sections generated with a cryostat (Leica) were then mounted (Prolong™ Gold Antifade Mountant, ThermoFisher). The interface of the chitosan film into the gel was assessed using confocal microscopy (LSM710, Zeiss). Samples were imaged using a confocal microscope for the tough hydrogel (nile blue, excitation 633 nm) and chitosan (FITC, excitation laser 488 nm).

Scanning Electron Microscopy.

Elemental mapping was performed on native (uncoated) samples under low vacuum conditions (20 Pa) using a Tescan (Brno, Czech Republic) Vega GMU scanning electron microscope equipped with a Bruker XFlash 5030 dual-detector EDS system. Each of the 512 × 512 elemental maps was acquired at an acceleration voltage of 20 keV and at a 15 mm working distance.

Evaluation of Zeta Potential.

Samples of chitosan (85 and 95% degree of deacetylation), acrylamide, alginate, and alginate-AAm were dissolved in water and/or HBSS. The resulting solutions were analyzed with a particle size analyzer (Malvern zen3600) for its zeta potential.

Burst Pressure Testing.

Standard tests (ASTM F2392) were completed for assessment of burst strength. A 15-mm diameter tough adhesive with and without attachment of a chitosan film to its back was evaluated. The pressure was applied by pumping air using a syringe pump (Harvard Apparatus PHD 2000 Dual Syringe Pump) at a rate of 2 mL/min through a 3 mm defect at the bottom of the sample. During measurement, the pressure was recorded, and the burst pressure was identified when a burst occurred.

For dynamic tests, a peristaltic pump was utilized. Aortic tissue (Sierra Medical) was carefully prepared and attached to liquid tight clamps. A 1 cm longitudinal defect was made and was repaired (4-0 Prolene) followed by application of the tough adhesive with or without a chitosan film. Pump rates (80 to 100 cycles/min) were completed, and the pressure was controlled with a clamp on the distal side.

Cooling Effects of Gels.

Tough hydrogels and TenderCare® samples (15 × 15 × 1.5 mm) were cut and placed on porcine skin for 1 min. The temperature before and after placement was recorded with an IR camera (10:1 Infrared Temp-Gun #2267-20, Milwaukee).

High-Frequency Ultrasound Imaging.

High-frequency ultrasound imaging (Visualsonics, VEVO 3100, 50 MHz transducer) was used to evaluate placement of the gels over bovine tendon and nerve tissues.

Supplementary Material

Appendix 01 (PDF)

Movie S1.

Instant, tough adhesion was achieved between hydrogels under standard peel test conditions.

Movie S2.

Attachment of the TG-CF to VHB produced instant strong adhesion (up to >4000J/m²).

Data, Materials, and Software Availability

All study data in the article and/or supporting information have been deposited in Harvard Dataverse (https://doi.org/10.7910/DVN/NOKRXQ) (35).