Adhesive hydrogels for bioelectronics

Benefiting from the unique advantages of superior biocompatibility, strong stability, good biodegradability, and adjustable mechanical properties, hydrogels have attracted extensive research interests in bioelectronics. However, due to the existence of an interface between hydrogels and human tissues, the transmission of electrical signals from the human tissues to the hydrogel electronic devices will be hindered. The adhesive hydrogels with adhesive properties can tightly combine with the human tissue, which can enhance the contact between the electronic devices and human tissues and reduce the contact resistance, thereby improving the performance of hydrogel electronic devices. In this review, we will discuss in detail the adhesion mechanism of adhesive hydrogels and elaborate on the design principles of adhesive hydrogels. After that, we will introduce some methods of performance evaluation for adhesive hydrogels. Finally, we will provide a perspective on the development of adhesive hydrogel bioelectronics.


Introduction
All cells in the human body can generate electrical signals, and the generated biological circuitry plays critical roles in the development, metabolism, regeneration, and physiological functions of the human body [1][2][3]. Many diseases, such as arthritis, irritable bowel syndrome and diabetes, are linked to impaired biological circuitry. Functional impairment also disrupts the flow of bioelectrical signals, which in turn leads to chronic systemic dysfunction. Moreover, in modern medicine, these electrical signals are an important basis for clinical diagnosis and disease analysis, which is very helpful for patients in surgery and intensive care [4][5][6]. Therefore, bioelectronic devices based on human body electrical signals show broad application prospects in health monitoring and analysis [7][8][9]. However, traditional electronic materials, such as silicon, platinum or platinum-iridium alloys, are rigid and cannot adapt well to the curved environment of human tissues or organs [10]. In addition, electrodes made of these rigid electronic materials could cause tissue damage when implanted into the human body, and the isolation space between the electrodes is particularly large, so the resolution of the detection signal would be low. Worst of all, the implanted electrodes would cause a reject reaction of the human body, and even produce a severe immune reaction, resulting in difficult healing of the surgical wound, and even endangering human life [11,12]. These shortcomings greatly limit the development and application of bioelectronic devices in the medical field. Therefore, it is necessary to develop flexible electrode materials with biocompatibility, high sensitivity and strong stability to meet the application of bioelectronic devices in the medical field.
Hydrogel is a kind of water-containing polymer material with the three-dimensional crosslinking network, which has superior biocompatibility, stable mechanical properties, good biodegradability, and adjustable mechanical properties, exhibiting unique advantages in bioelectronics [13,14]. Youn Soo Kim et al. prepared a conductive hydrogel by simultaneously initiating graphite stripping and polymerization of zwitterionic monomers by microwave radiation [15]. The conductive hydrogel closely resembles the outer neuronal electrodes of tissue and can meet the requirements of bioelectronics. In addition, it has been demonstrated that it can modulate rat sciatic nerves by low-current electrical stimulation. Guo et al. reported a double-network conducting polymer hydrogel composed of poly (3, 4-ethylenedioxythiophene): poly(styrenesulfonate) and polyvinyl alcohol (PVA) [16]. They adhered the conductive hydrogel to muscles for stable and long-term in vivo electromyography (EMG) recordings in a rat model. When the hydrogel adheres to the sciatic nerve, reliable electrical stimulation can be performed at stimulating voltages as low as 125 mV. The ideal hydrogels used in bioelectronics should possess high electrical conductivity, tissue-like mechanical properties, low toxicity, and minimal immune reactions. Most importantly, they should adhesion with biological tissues, which can enhance the contact between electronic devices and tissues, reduce contact resistance, and thus improve the performance of bioelectronic devices [17,18]. In order to achieve adhesion, chemical, and physical forces are usually introduced between hydrogels and biological tissues, including metal complexation, π-π superposition, cation-π interaction, covalent bonding, hydrogen bonding, and hydrophobic interaction [19][20][21]. In the following section, the mechanism of these interaction forces in adhesive hydrogels will be discussed detail, followed by elaboration on the design principles of adhesive hydrogels from the point view of the application, and this review will also introduce the methods of performance evaluation in adhesive hydrogels ( Figure 1). Finally, we will provide a perspective on the future development of adhesive hydrogel bioelectronics. We hope that this review will serve as a guide for the practical application of adhesive hydrogels in bioelectronics.

Mechanism of adhesion
The key to adhesive hydrogel is that a strong adhesion can be formed between the tissue and the hydrogel [22][23][24]. As shown in Figure 2, when the adhesive hydrogel is in close contact with the tissue, the adhesive layer is formed by connecting the surface functional groups of hydrogels with the tissue surface functional groups through chemical and physical interactions [25][26][27]. Many functional groups in hydrogels, such as hydroxyl, carboxyl, amino, ether, or catechol-group, can react with functional groups on the tissue surface to form imines, amides, or other covalent bonds [28,29]. Meanwhile, some non-covalent and physical interactions, such as hydrogen bonding, cation-π interaction, physical entanglement and interlocking, are also essential to adhesion [30][31][32]. Covalent bonds are typically strong, but they are irreversible. While noncovalent bonds are relatively weak, they are reversible, and they can allow for repeated adhesion and peeling [33][34][35].

Covalent bond
A covalent bond is a kind of chemical bond in which two or more atoms jointly use their outer electrons to achieve the state of electron saturation, thus forming a relatively stable chemical structure. Most hydrogels are rich in functional groups, which can further react with biological tissue surface functional groups to form covalent bonds at the interface between hydrogel and tissue [36][37][38]. In general, covalent bonds have high energy, so interface adhesion by covalent bonds is usually strong and stable. The common functional groups on the interface of biological tissues include the hydroxyl group, carbonyl group, carboxyl group, amino group, amide group, sulfhydryl group and guanidinium group [16]. Michael addition reaction and Schiff base reaction are common reactions to form covalent bonds ( Figures  3a and b). For example, hydrogels containing aldehyde groups can covalently bind to amino groups in biological tissues by Michael addition reaction, exhibiting high adhesion strength. In addition to aldehyde groups, the adhesive hydrogel reactive groups also include N-hydroxy succinimide esters, cyano groups, catechol and enzymes ( Figure 3c). The chemical structures and chemical reactions associated with these groups will be described in this section. Aldehyde groups. The structure of the aldehyde group is -CHO, which is a hydrophilic group. There is an electronegative imbalance between the oxygen and carbon atoms in the carbonyl group of the aldehyde group, so the aldehyde group has very high reactivity. The aldehyde can combine with amine groups to undergo a Schiff base reaction, forming imine bonds while releasing water molecules. This reaction is very fast at room temperature, taking only seconds to minutes. Most of the chemical bonds formed by the reaction with aldehyde groups are dynamic bonds of low strength. The kinetics of dynamic bond formation is usually associated with the pH of the surrounding environment, with the fastest rate of dynamic bond formation in weakly acidic environments [39]. Aldehyde groups are widely used because of their high reactivity and fast reaction speed at suitable pH environments. For example, Tariq Ahmad Mir et al. reacted synthetic guar gum (GG) monoaldehyde with glycerol used as a crosslinker to obtain a superabsorbent hydrogel (SAH) [40]. The hydrogel is capable of absorbing and retaining large amounts of water and artificial biological fluids without degradation and loss of physical stability, which enables it to be used as a transport vehicle for food nutrients. In addition, aldehyde groups can also act as a bridge to enhance the adhesion between the adhesive hydrogel and the biological tissue. Wu et al. successfully prepared an adhesive hydrogel using dopamine, carboxymethyl chitosan, and aldehyde chondroitin sulfate oxidized from chondroitin sulfate [41]. This hydrogel is an excellent candidate dressing for full-thickness wound healing with strong adhesive force, antibacterial properties, hemostatic properties, and spray film formation ability. Although aldehydes have the potential for adhesive hydrogels, the use of aldehydes often causes cytotoxicity and tissue toxicity. Therefore, the clinical application of aldehydes is limited. N-hydroxy succinimide ester group. N-hydroxy succinimide esters (NHS) are common reactive acylation reagents in biochemistry. They spontaneously react with primary amines under physiologically weak alkaline conditions (pH 7.2-9) to form amide bonds and release NHS leaving groups [39]. These reactions are relatively fast, ranging from seconds to minutes, which enables NHS esters to be widely used in adhesive hydrogels. For example, Nasim Annabi et al. constructed a robust adhesive hydrogel with painless detaching properties and in vitro biocompatibility by using N-hydroxy succinimide ester, polyethylene glycol diacrylate, sodium alginate, Fe 3+ and tannic acid [42]. It can not only produce rapid adhesion to wet tissue surfaces within 5s of gentle pressing, but also deform over time under wet conditions to support physiological tissue functions. Application of the hydrogel on rabbit conjunctiva and pig cornea showed that the hydrogel had strong adhesion. However, there are some disadvantages to NHS esters. They are highly reactive to nucleophiles, so they are easily hydrolyzed, and NHS esters should only be stored for long periods of time under dry conditions. Cyanide group. The cyano group is a group of carbon and nitrogen atoms linked by a triple bond. They are widely used in organic chemistry and can be used to synthesize many organic compounds, such as drugs, spices and dyes. However, due to the high toxicity of cyanide, it is necessary to use it with extreme care and take appropriate safety measures. Common cyano groups include cyanoacrylate, isocyanate and azide. The monomer of cyanoacrylate contains alkoxy group (-ROR), carbonyl group (-COOR) and cyanide group (-CN). These groups are highly electronegative and therefore can react with amine groups of tissues through a Michael addition [43]. Covalent cross-linking between polycyanoacrylates can be achieved by introducing amines from the tissue surface to the polymer chains as the Michael type initiators during hydrogel polymerization [44]. This process is usually completed within a few seconds, and therefore it can be used to synthesize adhesive hydrogels in situ. However, the skin application of cyanoacrylates is limited by their lack of degradability and toxicity.
Isocyanates are another type of reactive group providing −N=C=O−, which shows high reactivity to different nucleophiles because of the different electronegative properties between N, C and O atoms. When the isocyanate meets the primary amine, the nitrogen atom is coupled to the hydrogen atom of the primary amine, and the carbon atom is cross-linked to the nitrogen-hydrogen atom, forming a urea bond as the end substance of the reaction. For example, Emine Cansu Tarakci et al. synthesized a series of reactive hydrogels bearing isocyanate functional groups by photopolymerization of 2-isocya-natoethyl methacrylate (ICEMA) and poly (ethylene glycol) methyl ether methacrylate (PEGMEMA) [45]. These hydrogels can be easily functionalized with bioactive ligands, allowing ligand-directed protein immobilization.
Aryl azides can also be incorporated into tissue adhesive hydrogels because of their photoactivated azide units and benzene rings. Sumit Dadhwal et al. reported the synthesis of a bioorthogonal-responsive low molecular weight diphenylalanine (PhePhe)-based hydrogel that is capped with a 4-azido-2,3,5,6-tetrafluorobenzyl carbamate self-immolation linker [46]. The addition of tetrafluoride aryl azide groups enhanced the stability of the hydrogel in unbuffered water at lower critical gel concentrations, as well as enhanced the sensitivity to the bioorthogonal reagent trans-cyclooctene. Catechol. Catechol compounds have similar chemical properties to mussel adhesion proteins and are therefore widely used in adhesive hydrogels [30,47,48]. Catechol is a phenyldiol consisting of a benzene ring and two adjacent hydroxyl groups, which is a strong reducing agent and can react with a variety of functional groups ( Figure 3d). When catechol is oxidized to quinones, the benzoyl ring of the quinones can undergo a Michael addition reaction with the amine group on the tissue surface, and a Schiff base connection can be formed between the carbonyl group and the amine group. For example, Lu et al. fabricated a dual-network (DN) hydrogel based on bioadhesive catechol-chitosan hydrogel using polyacrylamide [49]. This DN hydrogel can be used for precision treatment and regulation of the inflammatory microenvironment of chronic wounds. Besides, Gao et al. prepared a kind of hydrogel with a synergistic strategy of combining catechol-Fe 3+ complexes [50]. The resulting hydrogels exhibited seamless self-repairing behavior, tissue adhesion and high mechanical property. Hossein Montazerian et al. proposed a simple oxidative polymerization step before the conjugation of catechol-carrying molecules (i.e., 3,4-dihydroxy-L-phenylalanine, L-DOPA) as a potential approach to amplify catechol function in adhesion of natural gelatin biomaterials [51]. Poly(L-DOPA) conjugates enhance the cross-linking density, reduce swelling and enhance cohesion, enabling stronger biological adhesion at body temperature. Enzyme. Enzymes can be used to promote cross-links within or between molecular chains to enhance the strength of adhesive hydrogels [52]. Without chemical cross-linking agents, the hydrogel formed by enzyme catalysis has the advantages of good biocompatibility, easy degradation, and low cytotoxicity. Therefore, enzymatic catalysis is also an environmentally friendly strategy for adhesion in hydrogels. For example, Lu et al. developed a mussel-like nanoenzyme using the natural polyphenol with ultrafine silver nanoparticles that catalyzes the hydrogel without additional conditions [53]. The mechanism of this enzyme is similar to that of mussel adhesion, which maintains the dynamic redox balance of phenol-quinone and endows the hydrogel with repeatable long-term adhesion, making it an adhesive bioelectrode for detecting of human physiological signals. The method of adding enzymes allows hydrogel cross-linking to occur generally at physiological pH and temperature, with high substrate requirements, but without the need for other radical initiators or chemical additives.

Noncovalent bond
Not only covalent bonds but also noncovalent bonds play an important role in adhesive hydrogels. Noncovalent interactions mainly include physical entanglement and interlocking, hydrogen bonding, and hydrophobic interactions. The adhesion strength of noncovalent interactions is not as strong as that of covalent interactions. However, non-covalent bonds can act together with covalent bonds to achieve strong adhesion between the adhesive hydrogel and the tissue [54,55]. The various noncovalent interactions will be described in detail in this section. Physical entanglement and interlocking. Polymer entanglement and mechanical interlocking are non-covalent bonding behaviors commonly used to adhesive hydrogels. From a microscopic perspective, both biological tissues and hydrogels can be viewed as microporous or nanoporous polymer networks composed of macromolecules with some reactive groups on the outside. The adhesive hydrogel is coated on the surface of biological tissue, and the different groups react with each other to form an interpermeable network. This penetration is driven by the diffusion effect of the polymer, so the hydrophilic and hydrophobic, molecular weight will have some influence [56,57]. Moreover, when physical entanglement is coupled with chemical cross-linking, they can produce synergistic effects. As shown in Figure 4a, Yuan et al. prepared a nanocomposite hydrogel by partial chemical crosslinking of  ethyl] dimethyl-(3-sulfopropyl) ammonium hydroxide (SBMA) and acrylamide (AM) monomer triggered, combined with physical crosslinking between cellulose nanocrystal and P(SBMA-co-AM) [58]. The hydrogel exhibits excellent adhesion, antibacterial properties and biocompatibility, and is able to detect external stimuli through ion signals and also mimics the excellent sensitivity of human skin. The specific method is to add polymer chains with polyfunctional groups and triggerable crosslinking properties to the adhesive interface. Polymer chains can diffuse into the hydrogel network, where they can be cross-linked. The polymer network formed in situ stitches the hydrogel like needle-stitching by topological winding with the existing network, without the need for other specific chemical designs of hydrogel. In addition, the combination of macrostructure and nanoparticles can also be applied to achieve the interlocking of tissues and adhesive hydrogels. For example, silica nanoparticles are used as a link between an adhesive hydrogel and a tissue interface. Severine Rose et al. showed that a strong and rapid adhesion between the hydrogel and the tissue can be achieved at room temperature by spraying nano-solution drops on the surface of the hydrogel and then touching another tissue [59]. This method was demonstrated by pressing hydrogel sheets with the same or different chemical properties or hardness together for approximately 30 seconds using different concentrations of silica nanoparticle solutions to achieve firm bonding. Hydrogen bonding. Hydrogen bonding can also be used for tissue adhesion hydrogels, which allows relatively rapid reversible binding between hydrogel and tissue. Hydrogen bonding is another intermolecular force except for the van der Waals force, which is an electrostatic interaction, and its strength is between the chemical bond and the van der Waals force. When hydrogen bonds are formed between molecules, the intermolecular force increases, the fluidity decreases, and the viscosity increases. Functional groups on the surface of the tissue can form hydrogen bonds with functional groups of the adhesive hydrogel, such as hydroxyl groups, carboxylic acids, and primary amines [60][61][62]. The binding force of a single hydrogen bond is relatively weak, but when multiple hydrogen bonds are formed in the system, the binding strength can be increased. Moreover, hydrogen bonds usually need to be bound to other groups to form a more strengthened and stable adhesion. For example, in polyacrylamide-alginate hydrogel, while the acrylamide double bonds polymerize, the amide bonds on the side chain of polymer form multiple hydrogen bonds with multiple hydroxyl groups in trehalose, which greatly improves the adhesion strength of hydrogel [63,64]. Electrostatic interactions in combination with complementary multiple hydrogen bonds can also be used to enhance the wet adhesion of hydrogels. For example, Wang et al. fabricated tough wet adhesion of hydrogen-bond-based hydrogel (PAAcVI hydrogel) using copolymerization of acrylic acid and 1-vinylimidazole in dimethyl sulfoxide followed by solvent exchange with water ( Figure 4b) [65]. PAAcVI hydrogel exhibited very strong adhesion to both wet and dry tissues. In addition, this hydrogel also showed strong long-term stable adhesion under water and various humid environments, and was able to deadhere and adjust its adhesion on demand. Hydrophobic effect. In addition to the above two interactions, the hydrophobic interactions can promote the adhesion between hydrogels and tissues [66][67][68]. It has been demonstrated that increasing the number of alkyl groups first results in enhanced adhesion, whereas adhesion strength instead decreases when more than eight carbon atoms are reached in the alkyl chain, suggesting the presence of optimal hydrophobicity for tissue adhesion [69]. Wu et al. developed an amphiphilic polyurethane hydrogel coating that achieved strong adhesion in water. When the coating that is composed of a mixture of amphiphilic polyurethane and water-soluble solvent was immersed in water, it exhibited strong adhesion due to the directional accumulation of hydrophobic chain segments along the substrate surface [70]. Meanwhile, the hydrophilic chain segments underwent physical entanglement to form a hydrogel coating, which makes the substrate superhydrophobic under water. As shown in Figure 4c, Guo et al. reported a hybrid hydrogel with ionic and hydrophobic cross-linking networks. The hybrid hydrogel can firmly adhere to various substrates such as glass, polypropylene, silica gel and wood with a maximum adhesion strength of 100 kPa [71]. Moreover, the hybrid hydrogel can stretch more than 8-10 times of initial length. The strong adhesion and high toughness property is attributed to the synergistic effect of electrostatic interaction and hydrophobic association. However, there will be some problems with hydrophobic interaction. For example, the binding strength of hydrophobic interactions is usually low. Fortunately, hydrophobic groups can act as repellents to replace water on the tissue surface. This prevents exposure of adhesion reactive molecules to water and enhances their reactivity. Such mechanisms are common in nature, including insect footpads and mussel secretions. Inspired by this mechanism, Fu et al. prepared a transparent, compliant, adhesive zwitterionic nanocomposite hydrogel. The zwitterionic polymers can form interchain dipole-dipole associations, providing an additional physical entanglement [72]. Reversible physical interactions endow hydrogels with rapid self-healing ability without any stimulation. Such hydrogels can adhere to many surfaces, including polyelectrolyte hydrogels, skin, eyeglasses, silicone rubber, and nitrile rubber. Chen et al. designed an ion-conducting gel for wearable underwater sensors through a fully hydrophobic structure [73]. This fully hydrophobic structure enables the ionic conductive gel to have strong underwater adhesion, superior water resistance, and excellent underwater stability.

Design principles for adhesive hydrogels
Besides adhesion properties, other important factors, including biocompatibility, mechanical matching, stress dissipation, fatigue resistance, swelling resistance, and degradability, and ease of removal also need to be considered when designing adhesive hydrogels [55,74,75]. It will be very difficult to satisfy all these factors in one material. Therefore, the researcher can first consider the required nature in conjunction with the clinical need. In this section, the design principles of adhesive hydrogels will be discussed from the perspective of chemical, biological, and physical characteristics.

Biocompatibility
The key role of bioelectronics is to generate and transmit electrical signals at the biological tissue-hydrogel interface, but there is a big premise-no large damage, inflammation and possibly cancer risk to the biological tissue when they can be used for the human body [76]. In other words, the adhesive hydrogel for bioelectronics must have good biocompatibility. The biocompatibility of adhesive hydrogels is mainly determined by their own chemical, biological, and physical properties [77,78]. Wang et al. developed supramolecular hydrogels with dynamic gel-sol-gel transition behavior driven by biocompatible cyclodextrin (CD) consumption [79]. As shown in Figures 5a and b, compared to the control samples, the hydrogel medium still showed high cell viability after 3 days of culture, and the cell density increased sharply with the increase of culture time. All these results indicate that the smart poly acrylic acid-C 18 (PAA-C 18 ) supramolecular hydrogel fueled with γ-CD is non-toxic, which is highly conducive to the development of biocompatible chemical fuel supramolecular materials that can be used in high-tech applications, especially in the field of bioelectronic medicine. For adhesive hydrogels, it may be necessary to use crosslinkers, initiators, accelerators in the synthesis process, and these chemical agents may produce by-products that cause inflammatory reactions. In addition, adhesive hydrogels based on natural polysaccharides, such as chitosan, may contain toxic substances and thus be toxic to humans if not properly purified. If natural materials are used as raw materials for adhesive hydrogels, the antigens of the natural materials themselves and the possibility of carrying pathogens should be considered [74,80]. The physical properties of adhesive hydrogels, including the structure, external end group, and molecular weight, also have a great impact on the biocompatibility of adhesive hydrogels. If the molecular weight of adhesive hydrogel is too small, once degrading in vivo, the by-products after decomposition may be harmful to the body. The high molecular weight of adhesive hydrogels may also cause renal stress response [81,82]. Besides, the hydrophobicity and hydrophilicity of adhesive hydrogel should also be considered. The surface free energy is closely related to the hydrophilicity and hydrophobicity of the adhesive hydrogel surface. And the hydrophilicity and hydrophobicity of adhesive hydrogels are the decisive factors for protein adsorption and cell adhesion. It has been proved that the surface of hydrogels with moderate hydrophilicity can promote cell growth [83]. The hydrophilic hydrogel surface can provide a moist and closed growth environment to promote cell proliferation and epithelial cell migration. However, excessive hydrophilicity of hydrogels may lead to excessive swelling, which may compress the tissues in contact and weaken the adhesion between hydrogels and tissues [84]. Therefore, the hydrophilicity of adhesive hydrogel needs to be within a certain range. The charge characteristics are also a factor that cannot be ignored when designing adhesive hydrogels. The above biocompatibility considerations determine the raw material of choice for the preparation of adhesive hydrogels.

Mechanical matching
A fundamental parameter for measuring the mechanical properties of adhesive hydrogels is the elastic modulus, which indicates the ability of the material to resist deformation when stress is applied. The lower elastic modulus, the better flexibility of the adhesive hydrogel. Different biological tissues have different ranges of elastic modulus values [85,86]. The mechanical difference between adhesive hydrogel and tissue easily generates mechanical stress concentration at the interface between tissue and adhesive hydrogel. This disrupts the chemical and physical connection of the tissue to the adhesive hydrogel interface and renders the adhesion at the interface ineffective [87,88]. Thus, when designing adhesive hydrogels, the mechanical matching between adhesive hydrogels and human tissues needs to be considered. Zhou et al. reported an organic hydrogel mechanically and functionally similar to human soft tissue by constructing a large number of sacrifice bonds (hydrogen bonds) between the soft hydrogel matrix and glycerol (Gly) through a solvent replacement strategy (Figure 5c and d) [89]. This organic hydrogel exhibits excellent mechanical properties completely matching with human soft tissue, such as a toughness of about 15.66 MJ m -3 , fracture  energy of 13.99 kJ m -2 , and Young's modulus of about 128 kPa. Using these advantages, this organic hydrogel has a wide range of applications, such as personal health monitoring, anti-counterfeits, and soft electronic skin, opening a way for the development of soft materials with complex mechanical properties and human-tissue-like functions. Moreover, if the adhesive hydrogel is too hard, it will hinder tissue movement, leading to tissue damage and inflammatory reaction. In serious cases, it will form scars. Therefore, the elastic modulus of the tissue should be carefully considered when designing adhesive hydrogels. The elastic modulus of the adhesive hydrogel can be adjusted by changing the molecular weight, and crosslink density of the polymer. The larger molecular weight of the hydrogel, the larger elastic modulus. The higher crosslinking density, the higher elastic modulus of the hydrogel. For example, using glycidyl methacrylate and methacrylic anhydride to modify regenerated silk fibroin (RSF) and gelatin, respectively, the mechanical properties, biocompatibility, and biodegradability of composite hydrogels can be adjusted by adjusting the cross-linking density of the polymer [90].

Stress dissipation
The hydrogels would undergo dynamic mechanical deformation (compression, bending, stretching, etc.) when in conformal contact with soft biological tissues. Therefore, the stress dissipation ability is an indispensable consideration when designing adhesive hydrogels. Since hydrogels have a gel state and good rheology, they have a low permeability coefficient, which can effectively decrease the internal friction of the medium, thus improving the dissipation energy. For example, Dominique P. Pioletti and his team used the original combined mechanism of flow-dependent and flow-independent dissipation sources to develop hydrogels with cartilage-like characteristics [91]. As shown in Figure 5e, macromolecular structures in such flexible networks can be reorganized to dissipate energy at deformation by rearrangement of polymer chains and reversible dissociation of physical bonds. By using a combination of permanent and temporary crosslinks, soft hydrogels consisting of lactose-modified chitosan have been developed. Although temporary cross-linking will have some effect on the shear modulus of the hydrogel, it will lead to rapid dissipation of the applied external force [92]. However, when the dissipation energy of hydrogel is small, its internal friction force will increase, and the flow of the medium will become unstable, leading to uneven flow change [74]. Many efforts have been made to improve the toughness of adhesive hydrogels by chemical modification or copolymerization modification. A common and effective strategy is to integrate polymer networks with different cross-linking mechanisms together to form a two-network structure. One of the networks is strongly cross-linked and the other is weakly cross-linked. The network with strong crosslinking properties provides the connectivity of the whole network, and the network with weak crosslinking properties provides the sacrifice bond for the hydrogel to prevent further crack propagation from destroying the hydrogel. For example, poly(N-hydroxyethyl acrylamide)-agar hydrogels are dual-network structures formed by covalently crosslinked N-hydroxyethyl acrylamide and agar [93]. When a crack occurs in the hydrogel, the poly(N-hydroxyethyl acrylamide) network is strained at the tip of the crack, transferring the stress to the agar network. The stress would destroy the relatively weak agar network structure by pulling apart the agar network, which would lead to the dissipation of crack energy.

Fatigue resistance
The communication between hydrogel and biological tissue is bidirectional. To some extent, biological tissue cells can be excited by applying electrical stimulation to the hydrogel; On the other hand, biological tissues undergo dynamic changes that are transmitted to the hydrogel, causing the hydrogel to generate fluctuations in electrical signals, for example, pulse measurement. The normal range of the pulse is consistent with the heartbeat, ranging from 60 to 140 beats/min. The pulse can be measured tens of thousands of times a day. Therefore, this puts a particularly high demand on the fatigue resistance of adhesive hydrogels. In the case of cyclic loading, adhesive hydrogels need to continuously maintain adhesion. Although mechanical stress applied to an adhesive hydrogel may not result in the adhesion failure of hydrogel to the tissue within several cycle loads, the failure may be caused by fatigue of hydrogel after multiple cycles [94][95][96]. Therefore, to ensure a stable adhesion between the hydrogels and tissues, the adhesive hydrogel must be fatigue resistant. When the applied load is less than the breaking strength, although the hydrogel will not fracture, tiny cracks will appear inside the hydrogel. During repeated loading cycles, the number of microcracks would gradually increase and expand, eventually leading to hydrogel failure. In addition, repetitive loading on the adhesive hydrogel can cause fatigue inside the adhesive hydrogel and at the tissue-hydrogel interface, which eventually leads to the failure of the adhesive hydrogel. Although the adhesive hydrogel has certain fatigue resistance properties, it will eventually undergo fatigue fracture due to repeated loading damage and depletion of stress dissipation from crack propagation [97,98]. The current inspiration for the design of adhesive hydrogels with fatigue resistance comes from human soft tissues, such as cartilage, ligaments, and tendons, which exhibit very superior fatigue resistance [99]. For example, inspired by the nanostructured interface between tendon, ligament, cartilage and bone, Liu et al. found that linking the ordered nanocrystal domains of synthetic hydrogels on engineered materials produces a fatigue-resistant adhesion (Figure 6a) [100]. Besides, salt solutions have been introduced into polymer networks to improve fatigue resistance. For example, polyacrylamide-CaCl 2 (PAM-CaCl 2 ) hydrogel showed no obvious interfacial crack propagation during load-unloading cycle even at 1000% strain [101]. Experimental results and theoretical calculations confirmed that the difference in hydration between salt and polymer chains leads to the steric limitation of polymer chains. Spatial confinement replaces the physical entanglement of polymer chains, which slides to dissipate energy and effectively avoids stress concentration and hysteresis, resulting in super-ductile hydrogels with fatigue resistance. However, it should be noted that the adhesion of the PAM-CaCl 2 hydrogel to biological tissues has not been prove.

Swelling resistance
Swelling is a common feature of most adhesive hydrogels, which reduces the mechanical properties and is detrimental to the adhesion of hydrogels and surrounding tissues. Adhesive hydrogel materials that commonly exhibit swelling include polyethylene glycol, acrylic acid, gelatin, and collagen [102,103]. The volume expansion generated by the adhesive hydrogel upon swelling would exert physical compression on the surrounding tissue. Sometimes the compressive strength can exceed the physical tolerance capacity of the tissue, leading to severe tissue damage. Moreover, swelling can also alter the mechanical properties of adhesive hydrogels, because the increased water content in adhesive hydrogels will reduce the polymer density [104].
Various methods, such as increasing the cross-linking density and the introduction of unique interactions (abundant hydrogen bonding, electrostatic repulsion and strong coordination association and hydrophobic aggregations) have been adopted to increase the swelling degree of adhesive hydrogels. The higher crosslink density can increase the mechanical swelling resistance. For instance, Feng et al. prepared an adhesive hydrogel with excellent swelling resistance through solvent conditioning to form hydrogen bond crosslinking with polymer chains and immersion of sodium citrate solution to induce ion crosslinking and salting-out, whose tensile strength retention rate after expansion equilibrium could reach 96% of the original value (Figure 6b) [105]. The hydrogel can be used for stable adhesion and wet adhesion of various substrates. Besides, Lan et al. reported a cellulose-skeleton based composite hydrogel [106]. This composite hydrogel has the advantage of high swelling resistance under the action of hydrogen bonding, making it an efficient wearable underwater sensor. Moreover, although the effect of environmental pH value is often neglected, the pH-dependent swelling behavior of hydrogels has been described in many kind of literature. For example, poly N-isopropyl acrylamide-carboxymethyl cellulose (PNIPAm-CMC), a semi-interpenetrating network hydrogel, showed the lowest swelling rate in the acidic environment [107].

Degradability
When adhesive hydrogels are used in vivo, the degradability of adhesive hydrogels appears particularly important, as degradable adhesive hydrogels can obviate the removal after surgery [108]. Generally, it is appropriate to design the degradation rate of adhesive hydrogels according to the working efficiency of bioelectronics. Too fast degradation rate of hydrogel will not only fail to play its role, but also increase the risk of tissue damage. If the degradation rate is too slow, the long-term existence of bioelectronics may damage biological tissues, leading to increased chances of tissue inflammation [109][110][111]. Therefore, in some cases, the adhesive hydrogel is designed to start degradation about 3 weeks after surgery, and complete degradation 3 months later. For example, Nathaniel S Hwang and his team prepared a hydrogel composed of hyaluronic acid and gelatin to provide extracellular matrix support to mimic natural brain tissue, improving the degree of cell infiltration. They tested the efficacy of the hydrogel by dual-channel fluorescence imaging, and applied image segmentation analysis to the treatment area where the hydrogel was implanted to provide precise visual information and data. The degraded hydrogel was followed for 21 days after hydrogel injection, and remnants were found in the intestine, indicating that the degraded hydrogel was cleared by the hepatobiliary (Figure 6c) [112].
Degradation is usually caused by the interaction of adhesive hydrogels with the surrounding environment. Hydrolysis is a common strategy to achieve the purpose of adhesive hydrogel degradation. The hydrolysis of polymers is caused by backbone breaks of polymers or by breaks of readily hydrolyzable chemical groups and /or chemical bonds, such as anhydride, ester group, thioester, amido bond and imine groups [109,113,114]. The rate of hydrolysis of the adhesive hydrogel ranges from a few hours to several months, according to the water uptake. The factors influencing the hydrolysis of hydrogels include chemical and physical properties (composition, crystallinity, porosity, hydrophilicity and hydrophobicity, skeleton structure and morphology) and environmental characteristics (pH and temperature) associated with the adhesive hydrogel. For example, PEG-polyester hydrogels are mainly degraded by surface corrosion after intraperitoneal injection, and the degradation of these hydrogels in the subdermal layer largely depends on the dissolution of the gel surface and the hydrolysis of the polyester segment [115]. The other factor should be noticed is the pH of the human body. The human body's pH environment ranges from an acidic environment to an alkaline environment, which can act as catalysts to promote the degradation of hydrogels [116,117]. It has been shown that the degradation rate of many adhesive hydrogels is affected by environmental pH [118]. For example, poly(ethylene glycol)-hydrazone covalent adaptive hydrogel is stable at neutral pH but degrades rapidly at acidic pH [119].
Enzymes are also able to influence the degradation properties of adhesive hydrogels. Enzymes promote hydrogel degradation by hydrolysis. Some natural protein polymers, including gelatin and collagen, can be degraded by specific enzymes such as gelatinase and collagenase [120]. Some synthetic polymers, such as polycaprolactone and polyurethane, can also be degraded by specific enzymes [121][122][123]. For example, PEG polypeptide hydrogens undergo enzymatic hydrolysis through protease, and finally degrade into non-toxic PEG and neutral amino acids. Neutral degradation products can reduce irritation to adjacent tissues in the body [115]. Besides, the degradation process catalyzed by one enzyme can be degraded more rapidly by the synergistic action of another or more enzymes.
In addition to the above-mentioned factors, the subsequent processing of hydrogel degradation products is also an important point. Degradation products may include cleavage chains, oligomers, and small molecule by-products, which are normally removed from the body by the kidneys. The extreme molecular weight capable of passing through the renal pathway is about 70 kDa [38,81]. In contrast, degradation products with molecular weights greater than 70 kDa must be excreted through other routes, such as the liver. Therefore, the toxicity of degradation products must also be considered [124,125].

Ease of removal
In addition to the biodegradable hydrogels mentioned in above section, there are also a number of hydrogels that can't be degradable or need to be adhered repeatedly in clinical practice. Therefore, ease of removal from tissues is also an important point to consider when designing adhesive hydrogels. A portion of adhesive hydrogel will form chemical cross-links with functional groups on the tissue surface, so there is usually some degree of adhesion on the tissue surface. Removal of the adhesion between the hydrogel and the tissue requires mechanical debridement of the tissue under anesthesia. This method is not only time-consuming, but also has a risk of damaging the tissue, which can cause physical and mental pain to the patient. Therefore, there is an urgent clinical need for adhesive hydrogels that can be removed with gentler and less invasive methods. At present, the removal methods of adhered hydrogels mainly include adjusting the temperature and pH, and adding other substances to dissolve the hydrogel. For example, Zeng et al. fabricated hydrogel motion sensors using acrylamide, stearoyl methacrylate C18, alginate and polyacrylic acid, and temperature-responsive polymers, F-127 and F-68 [126]. The transition from stable adhesion to on-demand separation between the adhesive hydrogel and the interface can be achieved by lowering the temperature of F-127/F-68 or adjusting the pH of alginate/polyacrylic acid from acidic to basic. These properties make the hydrogel an agent-sensing system that can not only detect human movements, but also seal tissue damage and measure water pressure with reliable response. Besides, Zhang et al. synthesized crosslinkable supermonomers to prepare an easily disassembled supramolecular hydrogel for wound dressing [127]. When exposed to the FDA-approved drug memantine, the supramolecular hydrogel dissolves itself and is easily removed (Figure 6d). These methods are effective, but the operation is cumbersome, and the conditions are also harsh, such as pH changes, which are harmful to human skin. Therefore, more methods with simple operation and mild conditions need to be explored to achieve on-demand removal of hydrogels in the future.

Performance evaluation of adhesive hydrogels
The properties of adhesive hydrogels can be evaluated by mechanical, chemical, and biological tests. Hydrogel bond strength and rupture pressure are usually determined by mechanical tests, and degradation and swelling properties are often assessed by chemical tests. Biological testing is performed to monitor the cytocompatibility, histocompatibility, the host immune response and the effect on tissue healing of adhesive hydrogels. This section will briefly review common methods used to evaluate the properties of adhesive hydrogels. These methods can qualitatively and quantitatively determine the properties of adhesive hydrogels.

Evaluation of mechanical properties
The basic measurement parameter of the mechanical properties of adhesive hydrogels is the bond strength or bond energy, which is usually obtained by peeling tests, tensile tests, and lap shear tests (Figure 7) [128][129][130]. The peeling test can measure the resistance of adhesive hydrogels when external forces are applied to peeling. Among the different types of peeling tests, the T-type peeling test, in which two partially adhesive layers are stretched at the adhesion interface, is the most commonly used (Figure 7a). It should be noted that loading rates have an effect on the peeling tests. Tensile tests determine the adhesion strength by pulling the adhesive hydrogel away from the tissue surface (Figure 7b). The experimental setup is relatively simple, but such tests often require accurate sample alignment and force, and the involvement of other stresses needs to be avoided to avoid interference. The ability of adhesive hydrogel to withstand planar stress can be evaluated by lap shear tests (Figure 7c). Clinically used adhesive hydrogels often suffer from shear stress in vivo. Therefore, a lap shear test is often adopted to obtain the mechanical properties of adhesive hydrogels in vivo. It is important to note that the measurement results of mechanical properties may be affected by the geometry of the adhesive hydrogel, such as the length, width, and thickness. In addition to the geometric factors, all measurement results are usually affected by the elongation rate because of the viscoelastic nature of adhesive hydrogels and tissues. Although the above methods facilitate the comparison of mechanical properties between adhesive hydrogels and can help to select the best adhesive hydrogel for a particular application, they usually cannot reflect the possible tissue characteristics and the specific situation of the patient. Thus, further evaluation with the aid of in vitro models for adhesive hydrogels is required.
The fatigue properties of the hydrogel can be obtained by applying cyclic loads and measuring the crack propagation during cycling. Fatigue tests are generally performed with 90°or T-type stripping tests. The force value of cyclic loading (F) should be less than the peel force (Fpeel), otherwise the hydrogel will be peeled off immediately. Then, the rate of crack propagation at the interface and the interfacial fatigue threshold are calculated by the differential of the crack size and the number of cycles. Generally, the higher the interfacial fatigue threshold, the stronger the fatigue resistance of adhesive hydrogels.

Evaluation of chemical properties
Any materials that enter the medical level must go through strict safety performance evaluation. Therefore, hydrogels that are applied to the field of medical engineering need to undergo a series of chemical performance tests, including degradation properties and swelling properties [131,132]. A general approach to evaluate the degradation performance of adhesive hydrogels is to place them in a physiological medium or solution containing biological enzymes [38,133]. After the hydrogel sample is put into the solution, keep observing the changes of the sample, such as volume, surface cracks, and record the changes in the sample within a certain time interval. By analyzing the recorded data, the degradation rate and degree of gel in different solutions can be obtained. Although the simple in vitro experiments can provide a rough estimate of the degradation performance of adhesive hydrogels, in fact, it does not fully reveal the complex processes occurring in vivo. Regular sampling and weighing of the implants after implantation in vivo can be performed for a more accurate assessment of the degradation properties of the hydrogels, but it requires a lot of experimental animals as samples. However, obtained samples may exhibit high variability, which may result in results dispatching with the facts. Therefore, it is generally not applicable for postoperative monitoring in humans [134]. To address these issues, researchers have developed real-time fluorescence imaging and magneto-elastic resonance sensors to monitor the degradation of adhesive hydrogels. Although these techniques allow us to characterize the degradation process in a non-invasive way, they have not been widely used because these techniques are too complex. In addition, mechanical tests can be performed to evaluate the degradation properties of adhesive hydrogels. Degradation can destroy the physical properties of adhesive hydrogels. Therefore, the change of mechanical properties of hydrogels over time can be used to indirectly indicate the degree of degradation. Swelling performance is also an important parameter for evaluating adhesive hydrogels. Due to the existence of the crosslinking network, hydrogels can absorb plenty of water, and the amount of water absorption is closely related to the crosslinking degree. The water absorption of adhesive hydrogel decreases as the degree of cross-linking increases. At present, the characterization method of expansion behavior is mainly to calculate the expansion rate by the weight or volume change of adhesive hydrogel [135]. Similar to the in vivo degradation test, the adhesive hydrogel can also be implanted into the body, and then taken out after a period of time to observe the expansion of adhesive hydrogel by measuring its volume or mass. Although more physiologically relevant data can be obtained with this approach, low-cost and time-saving in vitro assays often remain the preferred approach.

Biological performance evaluation
By culturing cells with adhesive hydrogels and monitoring cell viability, the biocompatibility of adhesive hydrogels can be assessed [136]. The International Organization for Standardization (ISO) has proposed a standardized bioassay method (ISO-10933) to standardize the commercialization and clinical translation of hydrogels. The ISO has issued 17 relevant standards with a serial number of 10993. At the same time, the evaluation methods of biocompatibility have also been standardized, including genotoxicity, carcinogenicity and reproductive toxicity experiments (10993-3), cytotoxicity experiments in vitro (10993-5), local reaction experiments after implantation (10993-6), stimulation and sensitization experiments (10993-10), systemic toxicity experiments (10993-11). These in vitro experiments can quickly detect the biological effects of adhesive hydrogels, and they are often used to assess the cytotoxicity of adhesive hydrogels.
Histological analysis is the standard for qualitative and quantitative measurement of the biocompatibility of tissue in vivo. There are many staining methods for histological evaluation, among which Hematoxylin Eosin (H&E) staining is the most widely used (Figure 8) [137]. The alkaline hematoxylin staining solution mainly stained the chromatin in the nucleus and the nucleic acid in the cytoplasm with violet blue, and acidic Eosin primarily colors the cytoplasm and components of the extracellular matrix red. This method can detect the inflammation extent of tissue caused by adhesive hydrogels. Masson's trichrome stain (MTS) is another commonly used staining technique [138]. The dyeing principle of this method is related to the size of the anion dye molecule and the penetration of tissue. The anion dye with a small molecular weight is easy to penetrate tissue with a dense structure, while anionic dyes with a high molecular weight can only enter tissues with a loose structure. It is an effective method to detect connective tissue and assess the degree of tissue fibrosis. Further biological assessment is usually carried out through animal experiments, mainly including in vivo mitogen detection, toxicity detection, skin allergic reaction, and tissue section observation.

Conclusion and prospect
Adhesive hydrogels have attracted much attention over the last few decades, and they have shown great application potential in human health monitoring and bioelectronics. In order to achieve adhesion, physical and chemical interactions, such as hydrogen bonds, hydrophobic interaction, metal complexation, π-π superposition, cation-π interaction and covalent bond, are introduced between hydrogels and biological tissues. The interfacial adhesion and mechanical similarity make it possible to be compatible between hydrogel and tissue, which is conducive to the high performance of bioelectronics. Although much effort has been made by researchers on adhesive hydrogels, there are still many challenges and unmet needs that motivate further research in this field.
First of all, although researchers have a preliminary understanding of the adhesion mechanism of hydrogels, there is still a lack of in-depth and systematic theoretical research on it. Researchers have made many explorations on the adhesion mechanism of hydrogel; however, no consensus has been reached yet. The popular viewpoint is that bonding of hydrogels and tissue surfaces requires both surface bonding and internal bonding. Cohesive molecular interactions determine the strength of the hydrogel, while the strength between the interfacial hydrogel and the tissue is determined by surface adhesion. When cohesion interaction and surface adhesion interaction reach balance, hydrogels can exhibit excellent adhesion performance.
Secondly, there is still some room for optimization in the performance of adhesive hydrogels, such as the repeated adhesion properties. At present, the researchers have developed an adhesive hydrogel that can be repeated 30 adhesion cycles on pig skin [139]. However, with the increase in adhesion times, the adhesion performance of hydrogel decreased significantly. Therefore, how to improve the adhesion times of adhesive hydrogel is the focus of future work. In addition, the adhesive ability between the adhesive hydrogel and tissue interface is not strong enough to meet the needs of clinical application.
And, there are few studies on adhesive hydrogelsbioelectronic devices currently, especially in vivo. Moreover, the degradation performance of adhesive hydrogels is a major concern. The ideal hydrogel can be degraded or transformed into harmless substances that can be discharged from the body within a period of time. Although many researchers have developed hydrogels with degradation capacity, their degradation kinetics have not been clearly elucidated.
Finally, the development of an emerging field often relies on constant trial-and-error methods or occasional serendipitous discoveries. It is worth our attention that the future development of adhesive hydrogel bioelectronics will be mainly based on rational guidance and design principles for understanding the interactions at the interface of biological tissue-hydrogel interactions. Although hydrogel bioelectronics has blossomed in the last decade, much remains unknown to be explored. The translation of adhesive hydrogels with ultra-high electrical and tough biomechanical properties into real functional bioelectronic devices remains a great challenge in the field.