Triple‐network‐based conductive polymer hydrogel for soft and elastic bioelectronic interfaces

Conductive polymer hydrogels have greatly improved the compatibility of electronic devices with biological tissues for human–machine interfacing. Hydrogels that possess low Young's modulus, low interfacial impedance, and high tensile properties facilitate high‐quality signal transmission across dynamic biointerfaces. Direct incorporation of elastomers with conductive polymers may result in undesirable mechanical and/or electrical performance. Here, a covalent cross‐linking network and an entanglement‐driven network with conductive poly(3,4‐ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS) have been combined. The triple‐network conductive hydrogel shows high stretchability (with fracture strain up to 900%), low impedance (down to 91.2 Ω·cm2), and reversible adhesion. Importantly, ultra‐low modulus (down to 6.3 kPa) and strain‐insensitive electrical/electrochemical performance were achieved, which provides a guarantee for low current stimulation. The material design will contribute to the progression of soft and conformal bioelectronic devices, and pave the way to future implantable electronics.

High conductivity is crucial for the low interfacial impedance of electrode devices based on conductive hydrogels.In recent years, researchers have found that the addition of dopants (e.g., dimethyl sulfoxide (DMSO), ethylene glycol (EG), sorbitol, malic acid, and glycerol [30][31][32][33][34] ), and/or post-processing treatments (e.g., soaking PEDOT:PSS sample in acids 35,36 ) can improve the conductivity.The original strong interaction in PED-OT:PSS is broken by hydrogen bonding, thereby forming physically cross-linked conductive channels through π-π stacking of PEDOT segments.Although the conductivities of pure PEDOT:PSS hydrogels are high, 30 the rigid polymer networks cannot dissipate external stress effectively.Therefore, these hydrogels typically exhibit poor mechanical durability, 37 which may lead to device failure in a dynamic-moving tissue environment.][40] But hydrogels combined with multiple tough networks generally show high Young's moduli, which mechanically mismatch with soft biological tissues (with Young's moduli lower than 30 kPa 41 ).Consequently, the development of soft hydrogels with low impedance and high stretchability is still challenging.
In this study, we adopt a triple-network (TN) strategy to combine all the desirable performances.First, a covalently cross-linked network (consisting of poly(ethylene glycol) diacrylate [PEGDA]) was introduced into the PEDOT:PSS hydrogel to improve stretchability.After being soaked in acetic acid (HOAc), PEDOT was enriched by phase separation with PSS, resulting in enhanced charge transport. 35,36By this way, a stretchable doublenetwork (DN) hydrogel with high electrical conductivity was obtained.Importantly, an entanglement-driven network was introduced into the DN hydrogel, which remarkably reduced Young's modulus (down to 6.3 kPa).We chose polyethylene glycol (PEG) as the third network because its hydroxyl groups and polar backbone can interact with the sulfonate groups of PSS to yield a better plasticization effect and elevate the conductivity of PEDOT:PSS, together with improved water-retention capacity.As a result, for the triple-network hydrogel with high-molecular-weight PEG (TN-H), the fracture strain reached up to 900%, with a low impedance of 91.2 Ω•cm 2 .Besides, the diffusive PEG chains also can make the hydrogels spontaneously entangle with the tissue, thereby achieving seamless adhesion onto tissue interfaces (with an adhesion energy of 40 J/m 2 ).To the best of our knowledge, TN-H hydrogel, compared with other PED-OT:PSS hydrogel conductors, shows an ultra-high stretchability and ultra-low modulus (Supporting Information: Table S1), which makes it a promising candidate for soft and living bioelectronics.Furthermore, TN-H hydrogel was applied to electrical stimulation of the sciatic nerve in a rabbit model.It produced the action potential amplitude at a low current (down to 0.6 mA), demonstrating a superior electrode-neural interface than a commercially available device based on metal electrodes.
Polystyrene-block-poly(ethylene-ranbutylene)block-polystyrene (SEBS: Tuftec TM H1062) was bought from Asahi Kasei Co., Ltd.All chemicals are analytical level and are used according to the receiving status.

| Synthesis of PEGDA
PEG (15 g, M W = 20 kDa) was fully dissolved in anhydrous dichloromethane (100 mL) in N 2 .After stirring for 30 min in the ice bath, acryloyl chloride (0.25 mL) and anhydrous dichloromethane mixed solution was added.The solution was stirred at room temperature for 48 h.The reaction solution was washed with saturated sodium bicarbonate solution and deionized water, and the organic phase solution was collected.The organic phase solution was dried with anhydrous sodium sulfate, filtered, and dried by rotary evaporation.Finally, the precipitate was vacuum dried for 24 h to obtain the final product poly(ethylene glycol) diacrylate (PEGDA).PEGDA was characterized by 1 H-NMR (Bruker Ascend 400 spectrometer) and Fourier transform infrared spectroscopy (Bruker Vertex 70).

| Preparation of TN-H hydrogel
PEGDA (0.3 g, M W = 20 kDa) was mixed with PED-OT:PSS solution (1.0 g) by stirring well.Then, DMSO (13 vol%) and LAP (1 mg, photoinitiator) were added and stirred at room temperature.The mixture was filtered through a nylon filter (1.0 μm).Next, the hydrogel presolution was transferred to a polytetrafluoroethylene (PTFE) mould and cross-linked by UV radiation (365 nm, 10 W power) for 10 min to ensure complete gelation.It should be noted that the dark color of PEDOT:PSS solution retards photo-polymerization by strong light absorption.Therefore, it is necessary to control the depth of the solution.The dried hydrogel was soaked in sufficient acetic acid solution for 2 h, washed with deionized water, and dried.Then, for TN-H or TN-L hydrogels, 1.0 mL of PEG (30 wt%) aqueous solution was absorbed by the redried hydrogels, followed by washing with deionized water and drying.At last, the dried hydrogels were soaked in 1.0 mL of deionized water until the water was completely absorbed.

| Characterization of morphology and composition
The hydrogel samples were freeze-dried and the morphology was obtained using a scanning electron microscope (SEM) (Hitachi SU8100).Then, the samples were prepared into films, and the infrared spectra were characterized using a Fourier transform infrared (FTIR) spectrometer (Bruker Vertex 70).Next, the ratio of PEDOT to PSS in the samples was characterized by X-ray photoelectron spectroscopy (XPS) (Thermo Scientific K-Alpha).The prepared hydrogels were immersed in a sufficient amount of 1× phosphate-buffered saline (PBS) solution, and the quality change of the hydrogel samples was monitored in real-time to complete the swelling tests.Finally, the morphology of the stretchable gold electrode (s-Au) electrode was observed using an optical microscope (Leica MC170 HD).

| Characterization of mechanical properties
The hydrogel samples were cut into cuboids (40 mm in length, 20 mm in width, and 1 mm in thickness) and the tensile properties were determined using a universal tensile testing machine (MTS Criterion Model C42.503) at a constant speed of 0.1 mm/s.Calculation of energy consumption efficiency: the ratio of the integrated area in the hysteresis loop under the loading curve.

| Characterization of electronic properties
The conductivity of the samples was tested by the fourprobe method (Keithley 4200).A universal tensile testing machine and a multimeter (Keithley 2636B) were used to measure the resistance change under strain by stretching the sample at a constant speed of 0.1 mm/s.

| Characterization of electrochemical properties
Two dry TN-H hydrogel electrodes (with an area of 0.1 cm 2 and a thickness of 60 μm) were prepared on the SEBS substrate, and one end of the hydrogel was connected to the copper wire electrodes with silver glue.The samples were soaked in PBS solution in advance.The electrochemical properties of the samples were determined using an electrochemical workstation (CHI660E; CH Instruments Inc.), in which a threeelectrode method was used: two hydrogel electrodes were used as the working electrode and the counter electrode, saturated silver/silver chloride (Ag/AgCl) electrode as a reference electrode.In electrochemical impedance spectroscopy (EIS) tests, the impedance frequency range was from 100 kHz to 1 Hz.The charge-storage capacity (CSC) of the samples was measured by cyclic voltammetry (the range of voltage: −0.5 to 0.5 V, scan rate 10 mV/s), and the charge injection capacity (CIC) of the samples was measured by chronoamperometry (the range of voltage: 0-0.1 V).Two s-Au electrodes (with an area of 0.1 cm 2 and a thickness of 100 nm) were also prepared on the SEBS substrate by heat-source evaporation.The SEBS substrate was spin-coated with SEBS solution (mixed 200 mg SEBS and 1 mL toluene) on the glass substrates at a speed of 1000 r/min.The electrochemical test methods of the s-Au electrode were the same as those of the TN-H electrode.
The CSC (mC/cm 2 ) of the electrode was calculated by the following equation 42 : where v is the sweep rate (mV/s); j c (u) is the cathodal current density (mA/cm 2 ); E a and E c are the anodic and cathodic potential limits (V); and U is the electrode potential (V).
The CIC (µC/cm 2 ) of the electrode was calculated by the following equation 43 : where i is the amplitude of the stimulation current (A); p w is the pulse width of the stimulation signal (µs); and A is the geometric surface area of the electrode (cm 2 ).

| Characterization of adhesion
The samples (40 mm in length, 20 mm in width, and 1 mm in thickness) were closely attached to each substrate for 2 h, and then the adhesion properties of the samples were measured using a universal testing machine.The adhesion energy of the samples was obtained by the 90°peel test method at a constant peel speed of 1 mm/s.PET film was used as the soft backing, and cyanoacrylate was used to glue the soft backing on top of the hydrogel.

| In vivo sciatic nerve stimulation
Pasting a mask on SEBS substrate, the above hydrogel pre-solution was spin-coated on it at a speed of 500 r/min and the mask was removed.After being cross-linked by UV radiation (365 nm, 10 W) for 10 min to ensure complete gelation.The uncross-linked solution was washed away in a large amount of deionized water to obtain a patterned hydrogel electrode.The dried hydrogel electrode was soaked in a sufficient amount of HOAc for 2 h, then washed with deionized water and dried.The redried hydrogel electrode was soaked in PEG (30%, M W = 20 kDa) aqueous solution for 48 h, then washed with a large amount of deionized water and dried to obtain a TN hydrogel electrode.The s-Au electrode was also prepared on the SEBS substrate using the same metal mask by heatsource evaporation (100 nm, 8 Å/s).New Zealand white rabbits were anesthetized with 2% isoflurane in balanced oxygen under sterile and body-temperature warming conditions.Nerve stimulation was tested around the sciatic nerve with TN-H electrode, s-Au electrode, and commercial metal electrode (Friend Ship Medical, PNB0.8×2/90).The needle-recording electrodes were inserted into the gastrocnemius muscle of the stimulated leg.Amplifier gain was set at 1000 with the output of the amplifier's bandpass filtered at 300 Hz to 10 kHz.The electrophysiological recording was conducted with the leg in a naturally relaxed position to prevent any restriction of movement.Stimuli were delivered in the form of constant current pulses with intensity from 0 to 3 V and every 1 s to gradually elicit nerve responses.Compound muscle action potentials (CMAPs) were continuously recorded before and after stimulation and responses were saved to data files.

| Animal experiment protocols
All animal tests were compliant with the standard guidelines approved by the Beijing Neurosurgical Institute Ethics Committee, and all animal surgeries were reviewed and approved by the Committee on Animal Care at Beijing Neurosurgical Institute (202104010).

| Preparation and characterization of the TN conductive stretchable hydrogel
We have adopted a synthetic strategy using a triplenetwork structure to construct a hydrogel with high stretchability, low impedance, and adhesion.PEDOT:PSS is chosen as the conductive polymer, with DMSO dopant to improve the conductivity.To enhance the tensile properties of the hydrogel, a high-molecular-weight PEGDA (20 kDa) was used as the second network because of its low cross-linking density. 44,45The in situ photo-cross-linked PEGDA interpenetrated with PED-OT:PSS, followed by HOAc-soaking and water-washing, giving a highly conductive and stretchable DN hydrogel.The third network is composed of a linear PEG (20 kDa).Since the hydrogel is a three-dimensional porous material (Supporting Information: Figure S3), the free long PEG chains could diffuse into the system and form entanglement with other networks.Note that the pore size of the cross-linked PEGDA network used in this paper matches the diameter of linear PEG. 46Such a physically cross-linked network greatly improves the tensile properties, 47 while interacting intermolecularly with the interfaces to give adhesion. 48Through the above steps, soft TN hydrogel with high stretchability, low impedance, and adhesion was prepared (TN-H, Figure 1A).In addition, we prepared two different types of control hydrogels to demonstrate the key role of the third PEG network.The first control hydrogel consisted of cross-linked PEGDA and PEDT:PSS (DN-P), which did not contain the PEG network.The second one was prepared by the incorporation of DN-P with lowmolecular-weight PEG (0.6 kDa, TN-L), instead of long PEG chains.
To verify whether the PEGDA and PEG were successfully introduced into the hydrogel, we performed FTIR spectroscopy and SEM.As shown in Figure 1B, the absorption peak of C ═ O (1720 cm −1 ) can be detected in both PEGDA powder and TN-H xerogel, and the absorption peak of C ═ C (1640 cm −1 ) is only found in PEGDA.It proves that the PEGDA network has successfully penetrated into the PEDOT:PSS matrix and been fully cross-linked.The SEM image shows that TN-H hydrogel exhibits a porous structure (Figure 1C).Compared with control hydrogels, TN-H hydrogel has lower porosity and smaller pore size, with interconnected fibers appearing in the pores (Supporting Information: Figure S3).In addition, DN-P and TN-L hydrogels show comparable swelling ratios, which are lower than that of TN-H hydrogel (Supporting Information: Figure S4).The reason may be that the long PEG in TN-H hydrogel, unlike the short PEG in TN-L hydrogel, shows stronger interactions with other polymer matrices and can hardly be exchanged with water during swelling.These results imply that long PEG chains successfully get entangled and such a physically cross-linked network is stable to the aqueous medium.Moreover, we performed XPS measurements of the TN-H xerogel and PEDOT:PSS to unravel the structure change in the presence of DMSO dopant and acid post-treatment (Figure 1D).The S (2p) single peak with a binding energy of 168.2 eV was assigned to PSS, and the S (2p) double peaks of 164.8 and 163.8 eV were assigned to PEDOT.Through integral calculation, it can be concluded that the ratio of PEDOT to PSS increased from 1:2.5 to 1:1.5, which leads to an increase in electrical conductivity.Besides, through the combination of UV-curing and anisotropic swelling, TN-H hydrogel can be patterned by stencil printing (Supporting Information: Figure S5), which extends its practicality in micropatterned devices for precise neuromodulation in vivo. 49

| Mechanical properties of the TN conductive stretchable hydrogel
The TN-H hydrogel exhibited remarkable mechanical ductility.It maintained homogeneous and intact under stretching and twisting (Figure 2A and Supporting Information: Figure S6).It could reach 900% elongation under the uniaxial tension (Figure 2B).Importantly, the addition of long PEG chains greatly reduced Young's modulus to 6.3 kPa (Figure 2C), which is five times lower than that of DN-P hydrogel and mechanically matched with many soft tissue interfaces.The fracture energy of TN-H is more than three times higher than that of TN-L (Figure 2D), suggesting that the presence of long PEG chains enables the hydrogel to withstand more mechanical energy.TN-H hydrogel showed a large hysteresis under high strains (≥100%) in loading-unloading cycle tests (Figure 2E).In contrast, DN-P hydrogel (that has no PEG chains) and TN-L hydrogel (that has short PEG chains) showed less hysteresis during loading-unlodaing cycles (Supporting Information: Figure S7).We calculate the ratio of the integrated area in the hysteresis loop under the loading-unloading curve as the efficiency of energy dissipation.The TN-H hydrogel was found to have the highest efficiency (50.2% ± 2.8%) (Figure 2F).Therefore, we can conclude that long PEG chains in TN-H hydrogel consume most of the energy via sacrificial bonds/interactions, allowing the hybrid network to dissipate energy in large deformation (Figure 2H). 50,51ompared with the other reported PEDOT:PSS hydrogels, TN-H hydrogel has high stretchability and ultra-low modulus simultaneously (Figure 2G and Supporting Information: Table S1).In most reported cases, covalent-cross-linked elastic polymers behave as the scaffold to improve the toughness.To achieve such a high stretchability, high polymer content is needed, which inevitably leads to high modulus and/or low conductivity.For our TN-H hydrogel, the addition of an entanglement-driven PEG network did not increase the elastic modulus, but greatly plasticized the composite, giving the ideal mechanical performance.

| Electrical properties of the TN conductive stretchable hydrogel
The conductivity of TN-H hydrogel could be elevated from 5.8 to 30 S/m by acid treatment (Supporting Information: Figure S8), determined by the four-probe method, and the conductivity of the xerogel could reach 118 S/m.Compared with the other soft and stretchable PEDOT:PSS hydrogels, the conductivity of TN-H hydrogel is outstanding (Supporting Information: Figure S9 and Table S1).Benefiting from the great ductility, TN-H hydrogel could be used as an excellent stretchable electrical conductor (Figure 3A).The resistance change of TN-H hydrogel was found to be negligible during stretching, especially under large deformation.For example, the normalized change in resistance (ΔR/R 0 ) is only 0.55 at 500% strain (Figure 3B).Taking the geometry variation into consideration, the conductive PEDOT:PSS network has hardly been destroyed.The reason should be that most of the energy was dissipated by straightening the PEGDA network in small deformation, during which resistance was dominated by geometry change.As strain increases, energy was dissipated by straightening the entangled PEGs.This process possibly accompanied with reorientation of PEDOT:PSS, because of their dipole and/or hydrogen-bonding interactions (Figure 2H).Consequently, the charge transport parallel to the stress became more efficient, and thus the resistance tended to be constant (under strain higher than 200%).Such an intrinsic stretchable conductor, which shows stable resistance under high strain, should facilitate electrical sensing in scenarios with large deformation.
We further carried out EIS for TN-H electrode in 1× PBS solution. 58The three-electrodes method was adopted, in which both working and counter electrodes were made of TN-H hydrogels.We also prepared s-Au with high electrical conductivity for comparison.S-Au exhibits good stretchability through crack engineering (Supporting Information: Figure S10), which has been widely used as a stretchable conductor for bioelectronics.The impedance of the TN-H electrode is lower than that of s-Au at all frequencies due to the high interfacial capacitance of PEDOT:PSS.Notably, the low-frequency unit-area impedance of the TN-H electrode (with the impedance of 91.2 Ω [1 cm 2 ] at  30,36,40,[52][53][54][55][56][57] are in Supporting Information: Table S1).(H) Diagram illustration of TN-H hydrogel deformation.PEDOT:PSS, poly (3,4-ethylenedioxythiophene):polystyrene.
1 kHz) was several orders of magnitude lower than that of s-Au (Figure 3C).After stretching, it increased by less than 0.4 times even under 100% strain (Figure 3D), which was observably lower than that of s-Au (increased by 6.2 times).In addition, we calculated the CSC and CIC of the electrodes, which are key parameters for evaluating the efficacy of electrical stimulation.As shown in Figure 3E,F, the CSC value and CIC value of TN-H reached 46.50 mC/cm 2 and 283 μC/cm 2 , respectively, which were much higher than those of s-Au (1.75 mC/cm 2 and 28 μC/cm 2 ).By monitoring the cyclic voltammetry curves of the electrodes under stretching (Figure 3G), the CSC value of TN-H reduced to 85% under 100% strain, superior to that of s-Au (with a 50% decrease).Figure 3H shows that external stress had little effect on the CIC of TN-H, while that of s-Au decreased significantly to 10%.These results suggest that the TN-H electrode has high mechanical stability which can provide a guarantee for effective electrical signal exchanges across dynamic bioelectronic interfaces. 59,60

| Adhesion behavior of the TN conductive stretchable hydrogel
Interfacial adhesion is essential to reduce the contact resistance between the hydrogel and biological tissue and to improve signal reliability. 61,62We measured the adhesion strength by standard 90°peel tests.It can be seen from Figure 4A that TN-H hydrogel has high adhesion energy, while DN-P and TN-L hydrogels show almost no adhesion, indicative of the essential role of long PEG chains.We examined the adhesion of TN-H hydrogel on different substrates (Figure 4B,C), and found that the adhesion on hard and dense surfaces not affected by the hydrophilicity of the materials.For instance, the adhesion strength of TN-H hydrogel on plastics (43.70 ± 1.29 J/m 2 ) is comparable to that on glass (45.68 ± 1.84 J/m 2 ).Therefore, we deduce that the PEG chains in TN-H can diffuse onto the gel-substrate interface and form unselective bonding via dipole-dipole interactions to provide adhesion (Figure 4F). 63,64Although water molecules will weaken such interaction, TN-H hydrogel shows moderate adhesion on moist surfaces (e.g., pig skin and muscle).It is possibly because of the microporous structure that allows the diffused PEG chains to penetrate and form entanglements at the gel-tissue interface.Notably, the adhesion energy of TN-H remains unchanged when subjected to repeated adhesion-detachment tests for 10 cycles (Figure 4D,E), and there is no hydrogel residue left on substrates after detachment (Supporting Information: Figure S11).Such reversible adhesion, ascribable to the dynamic noncovalent interface bonding, is beneficial to the retrieving of our electrode from soft bio-tissues with minimal damage.

| In vivo electrical signal recording and neural stimulation of the TN-H hydrogel
Regarding the objective measure of electrical conductivity of the TN-H electrode, a battery of electrophysiological experiments was performed based on a rabbit model (Figure 5A).Unlike commercial metal electrodes, the TN-H electrode (Figure 5B) and s-Au electrode could be wrapped around the sciatic nerve for neural stimulation, and CMAPs of gastrocnemius muscle were recorded simultaneously (Figure 5C and Supporting Information: Figure S12).TN-H electrode, s-Au electrode, and commercial metal electrode all successfully evoked CMAPs signal.Compared to the s-Au electrode and commercial metal electrode, the TN-H electrode conformed better and seamlessly adhered to soft tissues, thereby showing a substantially better electrophysiological parameter (Figure 5D,E).For the TN-H electrode, a low threshold current of 0.6 mA was sufficient to evoke just discernible CMAPs.In contrast, 0.9 mA was required for the s-Au electrode to evoke a detectable response (Figure 5F).The stimulation intensity gradually increased until there were no further increases in CMAPs amplitude, and the maximal stimulus currents were obtained.We observed the maximal stimulation intensity of the TN-H electrode was significantly lower than that of the s-Au electrode or commercial metal electrode (Figure 5G), indicating that the TN-H electrode could recruit a larger number of nerve fibers with lower stimulation current.Stimulation strength was increased further by 20% above the current resulting in maximum amplitude.The amplitudes of the TN-H electrode and commercial metal electrode under supramaximal stimulation presented higher than that of the s-Au electrode (Figure 5H).It also demonstrated a great capacity for nerve fiber recruitment.With these results, we demonstrated that our TN-H electrode provides an excellent electrode-nerve interface with favorable conductivity and low impedance.

| CONCLUSIONS
We report a triple-network conductive hydrogel with outstanding softness (modulus as low as 6.3 kPa), excellent stretchability (with fracture strain of 900%), low impedance (down to 91.2 Ω•cm 2 ), and reversible adhesion (adhesion energy of 40 J/m 2 ).The presence of an entanglement-driven PEG network plays a key role in realizing the ideal mechanical ductility.It enables the TN-H hydrogel to fit perfectly with soft organs and nerves in the body, together with strain-insensitive electrical resistance and electrochemical impedance.Benefiting from all these characteristics, the TN-H hydrogel was used to realize low-current neural stimulation, of which the performance is superior to that of commercial electrophysiological electrodes.Our TN-H hydrogel can be applied to construct more different types of soft bioelectronic devices because it is compliant with multiple solution processes for high-resolution patterning (e.g., stencil printing), which can facilitate seamless integration between electronics with soft tissues and/or soft-bodied creatures.

F I G U R E 2
Mechanical properties of TN-H hydrogel.(A) Photographs of TN-H hydrogel under stretching and/or twisting.(B) The stress-strain curves of the TN-H, TN-L, and DN-P hydrogels.(C) Young's moduli of the TN-H, TN-L, and DN-P hydrogels (n = 4).(D) Fracture energy of the TN-H, TN-L, and DN-P hydrogels (n = 4).(E) Successive loading-unloading cycles of the TN-H hydrogel under different strain amplitudes.(F) Efficiency of energy dissipation of the TN-H, TN-L, and DN-P hydrogels upon tensile loading and unloading (n = 4).(G) Comparison of Young's modulus-fracture strain of TN-H hydrogel with other PEDOT:PSS hydrogels (detailed data and references

F I G U R E 3
Electrochemical properties of TN-H hydrogel.(A) A circuit composed of a power supply, a LED bulb, and the TN-H hydrogel in the unstretched and stretched states.The change in brightness of the LED light was not obvious, indicating that the hydrogel has good stretchiness (DC voltage = 3 V).(B) Resistance change of the TN-H hydrogel under strain.(C) Impedances and phase angles of the TN-H hydrogel electrode and s-Au electrode.(D) Changes of the impedance of TN-H electrode and s-Au electrode under strain.(E) Current density-voltage curves of the TN-H electrode and s-Au electrode.(F) Charge injection capacity (CIC) tests of the TN-H electrode and s-Au electrode.(G, H) Changes of charge-storage capacity (CSC) and CIC of TN-H electrode and s-Au electrode under strain.

F
I G U R E 4 Adhesion properties of TN-H hydrogel.(A) Adhesion strength of TN-H, TN-L, and DN-P hydrogels on pig skin tested by standard 90°peeling.(B) Adhesion of TN-H hydrogel to various substrates (n = 4).(C) Photographs illustrating TN-H hydrogel's adhesion on a variety of surfaces, including skin, sandpaper, plastic, rubber, and glass.(D) The first, fifth, and tenth repeatable adhesion energies of TN-H hydrogel on pig skin.(E) Instant and repeatable adhesion of pig skin using TN-H hydrogel.The pig skin and TN-H hydrogel were adhered to and detached repetitively with a time interval between each repetitive test of 10 s (n = 4).(F) A schematic diagram of the adhesion mechanism between TN-H hydrogel and the substrate.