A stretchable, permeable, and biocompatible fiber-reinforced hybrid hydrogel electrode for highly stable electrophysiological signal recording

Abstract

The synergistic optimization of mechanical strength, skin-like elastic modulus, electrode-skin impedance, permeability, and biocompatibility remains a critical challenge in the deployment of flexible electrodes as a central component of noninvasive electrophysiological signal recording. Here, we propose a fiber-reinforced hybrid hydrogel (FRHH) electrode that integrates the conductivity of poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) and titanium carbide (Ti₃C₂T_x) with the mechanical resilience of styrene–ethylene–butylene–styrene (SEBS) fibers within a PVA hydrogel matrix. The FRHH electrode demonstrates remarkable stretchability, with a tensile strain reaching up to 1485%, coupled with moderate tackiness. It also shows low impedance at a frequency of 1000 Hz at the electrode-skin interface (2829.3 Ω), which is significantly lower than the impedance of commercial wet electrodes (6654.5 Ω) and dry electrodes (17 611.2 Ω). Furthermore, the FRHH electrode showed excellent biocompatibility in preliminary in vivo tests, allowing for continuous on-skin application for up to 12 hours without causing inflammation or allergic reaction. The electrode maintains conductivity and signal integrity under significant deformation, making it suitable for continuous and stable recording of electrocardiogram (ECG) and electromyogram (EMG) signals, even during physical activity. Additionally, the FRHH electrode shows promise in EMG-based gesture recognition and can recognize precise muscle activation patterns. The FRHH electrode holds promise for a wide range of applications, including continuous health monitoring, athletic performance tracking, and medical diagnostics, and could significantly contribute to advances in noninvasive and wearable healthcare technologies.

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1. Introduction

Electrophysiological signals, such as the electrocardiogram (ECG) and electromyogram (EMG), encapsulate an ocean of information regarding the physiological state of a human body.^1,2 Precise recording of these signals is prerequisite and crucial for disease diagnosis and treatment, motion analysis, and daily health monitoring.^3,4 The gold standard for short-time electrophysiological signal recording is the Ag/AgCl gel electrode (so-called wet electrode) due to its low impedance and excellent corrosion resistance.⁵ While wet electrodes have favorable mechanical properties, the elastic modulus mismatch between Ag/AgCl and skin poses considerable challenges for applications requiring prolonged skin contact.^6,7 Over time, the stability of the electrode-skin interface can become significantly compromised, degrading the quality of the acquired signals.^8,9

To address this issue, flexible and stretchable electrodes have been proposed. These electrodes are typically achieved by affixing a conductive material such as liquid metal,¹⁰ conductive fibers,¹¹ carbon nanotubes,¹² graphene,¹³ or MXene^14,15 and their composites¹⁶ as the functional material atop a flexible substrate such as polydimethylsiloxane (PDMS) or Ecoflex, which can then be molded into versatile geometries.^17–22 While these approaches can mitigate the interfacial adhesion issues between the electrode and skin to some extent, they frequently exhibit poor permeability, which could potentially lead to local skin inflammation or allergic reactions, hindering their reliable long-term usability in on-skin electronic applications.^23,24 Thus, the development of a novel electrode with skin-conforming flexibility and permeability is essential for the efficacious collection and analysis of electrophysiological signals in the field of wearable electronics.

Reducing the electrode-skin interfacial impedance is another key factor in the accurate recording of electrophysiological signals. Flexible electrodes of wearable electronics intended for direct skin contact frequently encounter high impedance presented by dry skin, which poses a significant challenge.^25,26 With dry electrode systems, a higher pressure must be applied to the electrode to reduce the contact impedance between the electrode and skin.²⁷ This may cause significant discomfort and may even lead to skin damage. On the other hand, with wet electrode systems, such as those based on the Ag/AgCl electrodes, a conductive gel must be applied to reduce the electrode-skin impedance, significantly increasing the complexity of electrode usage. Moreover, the conductive gel gradually dries out over time, causing a rapid increase in electrode-skin impedance, which hinders its long-term use.²⁸ Recent research has focused on conductive hydrogel-based semi-dry electrodes (HBSDE).^29–31 By slowly releasing electrolytes, the electrodes can gradually infiltrate the stratum corneum, the outermost layer of the epidermis, which is primarily responsible for the high impedance of dry skin, helping to lower the overall skin impedance.

Although hydrogels can effectively improve the electrical properties of the electrode-skin contact, their low mechanical resilience is frequently cited as a limitation. The mechanical fragility of hydrogel substrates, characterized by a relatively low tensile modulus and poor extensibility, can lead to delamination or rupture during motion, impacting the signal-noise-ratio (SNR) and stability. Various methods have been recently proposed for creating high-strength hydrogels (HSH),³² including double network/multi-network configurations,³³ topological designs,³⁴ and supramolecular approaches.³⁵ While HSH may address the issue of material fracture, the resulting mismatch between the high modulus of HSH and the elastic modulus of skin results in an unstable electrode-skin contact, potentially leading to delamination and reducing reliability of the device. To improve the strength of HBSDE while reducing the degree of mismatch between the Young's modulus of the electrode and that of the skin modulus, one highly promising strategy involves incorporating high modulus and highly elastic skeleton materials into the hydrogel to share the stress during electrode deformation, without affecting the low modulus of the hydrogel in contact with the skin.

In this work, we report a stretchable fiber-reinforced hybrid hydrogel (FRHH) electrode for highly stable noninvasive electrophysiological signal recording. The FRHH electrode is composed of a flexible and permeable styrene–ethylene–butylene–styrene (SEBS) fiber substrate, a polyvinyl alcohol (PVA) hydrogel, and a bi-dimensional conductive material comprised of one-dimensional poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) and two-dimensional titanium carbide (Ti₃C₂T_x). This hybrid design leverages the reinforcement and stress concentration effects of the fibers, enabling the FRHH electrode to retain its structural integrity during up to 800% elongation and repeated deformation. The FRHH electrode demonstrates low electrode-skin impedance (2829.3 Ω at 1000 Hz) and excellent biocompatibility, presenting an effective approach for acquiring ECG and EMG signals in both resting and motion states, as well as performing EMG-based gesture recognition.

2. Experimental section

2.1. Materials

PVA (molecular weight = 1750 ± 50 g mol⁻¹, hydrolysis degree ≥99%) was purchased from Shanghai Xianding Biotechnology Co., Ltd. Polystyrene-block-poly(ethylene-ran-butylene)-block-polystyrene (SEBS, average M_w ∼118 000) was purchased from Sigma-Aldrich. Sodium chloride (NaCl, 99.8%) was purchased from Shanghai Macklin Biochemical Co., Ltd. Tetrahydrofuran (THF, >99%) was purchased from Shanghai Aladdin Bio-Chem Technology Co., Ltd. PEDOT:PSS/deionized water conductive polymer (1.5 wt%) with 700–800 S cm⁻¹ conductivity was purchased from Shanghai Ouyi Organic Optoelectronic Materials Co., Ltd. Titanium(II) hydride powder (TiH₂, 400 mesh, 99%), Al (5–6 mm, 99.8%), and TiC (400 mesh, 98%) were purchased from Shanghai Aladdin Bio-Chem Technology Co. Ltd. Fluorocarbon surfactant (FS-3100) was purchased from DuPont Research & Development and Management Co. Ltd. All reagents were used without further purification.

2.2. Characterization

Micromorphological analyses were performed by means of field emission scanning electron microscopy (FE-SEM, LEO-1530, Zeiss, Germany) and transmission electron microscopy (TEM, Talos F200S G2, USA). The crystal structures of the samples were identified using X-ray diffraction (XRD, Bruker D8 ADVANCE, Germany) with Cu Kα radiation at a scanning rate of 8° min⁻¹ over the 2θ range from 5° to 85 °C. The elemental composition of the FRHH was analyzed using X-ray photoelectron spectroscopy (XPS, K-Alpha, Thermo Fisher Scientific, USA). Thermogravimetric analysis (TGA) was performed using a simultaneous thermal analyzer (TG-DSC, METTLER 1100LF, Switzerland). Stress–strain testing was conducted using a universal tensile testing machine (YL-S20, Guangdong Yuelian Instruments Co., Ltd, China). The element concentration in the FRHH was measured by inductively coupled plasma optical emission spectroscopy (ICP-OES, Agilent 5110, USA).

2.3. Preparation of the SEBS fiber film

First, SEBS was dispersed in THF at a ratio of 1 : 7 and stirred at 70 °C for 30 minutes. Then, the solution was loaded into a 2 mL syringe with a needle. During electrospinning, the injection speed of the solution was kept at a value of 1.0 mL h⁻¹ and positive 13 kV voltage was applied to the needle. The SEBS fiber film was collected using a metal plate which was applied negative 6 kV voltage.

2.4. Synthesis of Ti₃AlC₂ powder

Typically, TiH₂, Al, and TiC powders were mixed in a 1 : 1.2 : 2 mole ratio, followed by ball-milling for 18 h. The mixture was then sintered in a tube oven at 1450 °C for 2 h in a flowing Ar atmosphere. On crushing and sieving the resulting compact powder, Ti₃AlC₂ powder with a size of <38 μm was obtained.

2.5. Preparation of Ti₃C₂T_x suspension

The few-layer Ti₃C₂T_x suspension was synthesized by etching the Ti₃AlC₂ MAX phase in LiF/HCl hybrid etchant according to the previously reported route. Typically, a total of 3.0 g of Ti₃AlC₂ powder was gradually added (over the course of 3 min) to the LiF (4.8 g)/HCl (60 mL, 9 M) mixture. Then, the mixture was left under continuous stirring at room temperature for 24 h. Afterward, the resultant acidic mixture was washed repeatedly with deionized water via centrifugation at 3500 rpm for 5 min per cycle until the pH of the supernatant reached 6. The stable darkgreen supernatant was collected to obtain the few-layer Ti₃C₂T_x suspension with a concentration of ∼2.0 × 10⁻³ g mL⁻¹, which was determined by filtering a known volume of the suspension and measuring the weight of the film after vacuum drying.

2.6. Preparation of the FRHH

The precursor solution was prepared via the following procedures. First, 0.24 g NaCl was dissolved in 7.2 g deionized water and magnetically stirred at ambient temperature for 5 minutes, 0.8 g PVA was add to the NaCl solution, the concentration of the PVA solution for preparing the hydrogel was 10 wt%. The mixed suspension is stirred at 60 °C for 45 minutes at a slow rate to swell the PVA, and then at 98 °C for 90 minutes at a high rate to dissolve the PVA. The PVA solution, PEDOT/PSS solution, and Ti₃C₂T_x suspension are mixed in a mass ratio of 8 : 4 : 3 and stirred at ambient temperature for 1 h, yielding black mixed solution for preparing hydrogel.

The precursor solution was cast into the SEBS fibers covered with square container uniformly (the area of the square container is 4 cm × 4 cm). The square container is placed in the freezer for at least 4 hours before thawing, then repeat the same steps three times to complete the hydrogel crosslinking.

2.7. Animal experiments

All animal experiments were conducted in accordance with the relevant laws, regulations, and guidelines of the Laboratory Animal Ethics Committee of Guangxi University. Biocompatibility test was performed on 3 mice (C57BL/6), including 3 experimental groups for 2, 6 and 12 hours of attachment. First, mice were anesthetized with pentobarbital sodium (40 mg kg⁻¹) then put on 37.5–38 °C heating blanket for anesthesia observation. Second, commercial electrodes, dry electrodes, and FRHH electrode were attached to the dorsal skin of the test mice. After that, all mice were kept in three plastic cages with access to food and water ad libitum. The dorsal skins of mice were cut off after 2, 6 and 12 hours, respectively, and 4% perfluoroalkoxy (PFA). Hematoxylin–eosin staining was finally performed for pathologic analysis.

2.8. Skin impedance and human body electrical signal acquisition

All non-invasive human subjects' experiments were conducted in accordance with the relevant laws, regulations, and guidelines of the Ethics Committee of Guangxi University, and the informed consent was obtained from all human subjects involved in this study. The commercial electrode and dry electrode used for comparison are Monitoring Electrode with Foam Tape (2228, 3M health care). The skin impedance was measured by an electrochemical workstation (Metrohm Multi Autlab M204, Switzerland). The ECG signals were collected using Wireless acquisition system (KS108X-WEDB, Kingsense electronics, China), combined with the supporting Android program to realize real-time signal display and data preservation. The sampling rate was set to 500 sps. The EMG signals collected using Electromyographic System (Shenzhen Runyi Taiyi Technology Co., Ltd, China). The multi-channel signal acquisition was recorded by RTLAB software. The sampling rate was set to 1000 sps.

3. Results and discussion

3.1. Fabrication of FRHH electrode

To fabricate the FRHH electrode, SEBS was first dissolved in THF, forming an electrospinning precursor. After electrospinning and drying, a stretchable SEBS fiber film was obtained (Fig. 1a). FESEM and transmission TEM were used to analyze the micromorphology of the fiber film. As-prepared SEBS exhibited a uniform fibrous-network structure with an average fiber diameter of ∼10 μm (Fig. 1b and c). Subsequently, Ti₃AlC₂ was etched using a LiF/HCl solution to produce a layered Ti₃C₂T_x structure and dispersed in solution, as we previously reported (Fig. S1 and S2†).³⁶ Then the PEDOT:PSS solution and the etched Ti₃C₂T_x were dispersed and dissolved into a mixture consisting of PVA and 0.9% NaCl solution to form the hydrogel precursor. The precursor was blade coated onto the surface of the SEBS fiber substrate, ensuring a uniform distribution of the materials which reduces stress concentrations during deformation and results in a uniform conductive layer (Fig. 1d). Then, the compound was placed on a mold and underwent a cyclic freeze–thaw process (Fig. S3†). The freeze–thaw cycle represents a physical cross-linking approach for PVA hydrogels. During freezing, water crystallization excludes PVA chains from ice-forming regions, concentrating them in the non-frozen phase and promoting intermolecular hydrogen bonding between hydroxyl groups. Upon thawing, the ice melts while the hydrogen-bonded network remains, forming a cross-linked hydrogel structure.³⁷ In this study, PVA hydrogels were cross-linked through freeze–thaw cycles at −20 °C (freezing) and room temperature, respectively.³⁸ After that, a highly stretchable FRHH electrode with 57.1 wt% SEBS fiber substrate was obtained, and was easily removed from the mold and transferred for later use. SEM images revealed the detailed microstructure of the hybrid hydrogel, showcasing a distinctly porous architecture (Fig. 1e and f). By using a patterned mask plate, the fabrication process facilitates fabrication of FRHH electrodes with various patterns, thus allowing the electrode to be adapted to various application scenarios. The FRHH electrode with a thickness of 2 mm exhibited appropriate skin adhesion, and exceptional mechanical and electrical stabilities under different configurations such as bending and twisting (Fig. 1g and h), making it a promising candidate for application in wearable electronics and on-skin bioelectronics.

Fig. 1 Fabrication and preliminary characterization of the FRHH electrode. (a) Electrospinning process for preparing the SEBS fiber substrate; (b) SEM image of the SEBS fiber substrate and (c) a magnified image taken on the region marked by the red dashed box; (d) coating and crosslinking of the hybrid hydrogel precursor with the SEBS fiber substrate; (e) SEM image of the FRHH electrode surface and (f) a magnified image taken on the region marked by the red dashed box; (g) demonstration of the adhesiveness and stretchability of the FRHH electrode; (h) illumination of three LEDs in the initial, bending, and twisting states using the FRHH electrode.

3.2. Micromorphological and component interaction analysis of FRHH electrode

The integration of the SEBS fibers with the hybrid hydrogel was first characterized qualitatively using SEM analysis of sample cross-sections. As shown in Fig. 2a, the hybrid hydrogel was able to infiltrate the SEBS fiber substrate from the surface to the inside but did not completely permeate the substrate, instead forming a layered structure with a total thickness of 50 μm. The distribution of the conductive materials include Ti₃C₂T_x and PEDOT:PSS within the FRHH was comprehensively characterized using SEM and energy-dispersive spectroscopy (EDS). The SEM images reveal a two-dimension structure (Fig. 2b and c). The elemental mapping of sulfur and titanium in FRHH were also measured to confirm the distribution characteristics of Ti₃C₂T_x and PEDOT:PSS (Fig. 2d). The results show that Ti₃C₂T_x and PEDOT:PSS are both uniformly dispersed in the hydrogel matrix as electronic conductive media. This uniform distribution of conductive materials is instrumental in establishing continuous and interconnected conductive pathways throughout the FRHH, thereby significantly enhancing its electrical conductivity and stability. XRD was employed to assess the crystallographic structure of the incorporated materials (Fig. 2e). Characteristic diffraction peaks confirm the crystallinity of the Ti₃C₂T_x, SEBS, and FRHH electrode, demonstrating successful amalgamation of Ti₃C₂T_x into the SEBS matrix without compromising the structural integrity of either constituent. TGA indicated a mass fraction of SEBS fibers of 57.1%, a mass fraction of PVA of 19.5%, and the remaining mass is that of the conductive dielectric (Fig. S4†). Surface elemental composition was analyzed using XPS, and the results are presented in Fig. S5† and 2f–h. The spectra confirm an elemental composition predominantly comprising carbon, oxygen, and sulfur (Fig. S5†). High-resolution spectra show the presence of various functional groups. The C 1s spectrum reveals the existence of C–C, C–O, and C=O bonds, indicating a complex carbonaceous structure within the FRHH (Fig. 2f).³⁹ The O 1s spectrum is dominated by oxygen in hydroxyl (O–H) and carbonyl (C=O) groups (Fig. 2g), while the S 2p spectrum shows peaks corresponding to different sulfur oxidation states, likely due to the sulfonate groups in the PEDOT:PSS component (Fig. 2h). The schematic representation in Fig. 2i depicts the intermolecular interactions within the FRHH. In particular, hydrogen bonds and van der Waals forces between the hydrogel components (PVA, PEDOT:PSS, and Ti₃C₂T_x) and the SEBS copolymer with their polar and non-polar segments, respectively, are indicated. Intermolecular interactions, along with an extensive network of long-chain molecular structures within the FRHH, constitute the essential underpinnings of its mechanical robustness and flexibility. Although relatively weak chemical bonds, hydrogen bonds play a crucial role in intermolecular interactions in biology and materials science. The FRHH electrode has a multitude of sites available for the formation of hydrogen bonds, enabling it to interact with the skin surface through hydrogen bonding and exhibit moderate tackiness. This design strategy leverages the advantages of hydrogen bonding to enhance the contact between the electrode and the skin, which contributes to the improved stability and quality of the electrophysiological signal acquisition. In addition, the presence of sodium ions, highly dispersed Ti₃C₂T_x, and PSS groups endow the FRHH electrode with uniform conductivity (Fig. S6†).

Fig. 2 Microscopic morphology and elemental analysis of the FRHH electrode. (a) Cross-sectional SEM image of the FRHH electrode; (b) SEM image of the FRHH matrix substrate and (c) a magnified image taken on the region marked by the red dashed box; (d) SEM image of FRHH and corresponding element mapping of sulfur and titanium elements. (e) XRD patterns of FRHH electrode, Ti₃C₂T_x, and SEBS fiber substrate; the high-resolution spectra of (f) C 1s, (g) O 1s, and (h) S 2p orbital deconvolutions; (i) schematic diagram of the binding modes of FRHH electrode components.

3.3. Mechanical and electrical properties of FRHH electrode, and finite element analysis

Tensile testing was conducted using a universal testing machine, and the results reveal the excellent mechanical resilience of the FRHH electrode. Fig. 3a presents the stress–strain curves of hydrogels (2 cm × 1 cm) with different compositions. Notably, the tensile strength of hybrid hydrogels lacking individual conductive components shows no significant change. The fracture strain measured using an FRHH with larger length-width values (3 cm × 1.7 cm) can even reach 1485% (Fig. S7†). This indicates that the conductive components have no notable impact on the tensile strength. On the contrary, the FRHH demonstrates enhanced overall performance, further validating the fiber-reinforcing effect of the SEBS fibers. To explore the influence of various conductive components on the properties of hydrogels, the water absorption and conductivity of the hydrogel lacking different single components were measured (Fig. S8†). The results show that each component has no significant effect on the water absorption of the hydrogel. The conductivity of the experimental group without SEBS remained unchanged, because SEBS, as a mechanical substrate, did not affect the conductivity of the hydrogel, while the conductivity of the experimental group without PEDOT:PSS and Ti₃C₂T_x decreased because the electronic conductive path was reduced. The decrease in the experimental group lacking NaCl is due to the loss of ion conductive pathways. The AC impedance of hydrogels lacking different single conductive components was measured by using electrochemical impedance spectroscopy. The results showed that the AC impedance of the hydrogel was significantly reduced and the conductivity was significantly improved after adding various conductive components; for the components of NaCl lacking in ion conductivity and PEDOT: PSS/Ti₃C₂T_x lacking in electron conductivity, low impedance was observed at low and high frequencies. Therefore, the hydrogels lacking NaCl (electronic-conduction-only) and those without PEDOT:PSS/Ti₃C₂T_x (ionic-conduction-only) exhibit respective advantages in specific biosignal frequency ranges, in the ECG frequency range (0.25–35 Hz, low-frequency regime), the ionic conduction mechanism (via NaCl) dominates, whereas in the EMG frequency range (20–250 Hz, high-frequency regime), the electronic conduction mechanism (via PEDOT:PSS/Ti₃C₂T_x) becomes predominant (Fig. 3b).

Fig. 3 Mechanical and electrical properties of the FRHH electrode. (a) Fracture strain test of FRHH electrode and its various component combinations; (b) AC impedance test of FRHH electrodes and its various component combinations; (c) illumination of an LED using the FRHH electrode at 500% elongation, with brightness comparable to that at the original length; (d) resistance strain curve and GF strain curve of FRHH electrode; (e) AC impedance test of FRHH electrode after 3000 stretching–releasing cycles; (f) AC impedance values of FRHH electrode at 1000 Hz frequency after 3000 stretching release cycles; (g) cyclic tensile force–cycle number curves of FRHH electrode under 100% strain for 100 cycles and (h) the presentation of cyclic data at the initial, intermediate, and terminal stages of the experiment. (i) SEM images of the FRHH electrode before and after stretching and during the stretching process; (j) illumination of an LED under various configurations of the FRHH electrode; (k) the FRHH electrode retains excellent conductivity after being cut and reconnected; (l) the FRHH electrode remains stable and can effectively illuminate the LED in an underwater environment.

The ability of the FRHH to withstand extensive strain suggests superior toughness and suitability for applications requiring both durability and flexibility. The mechanical and electrical stability of the FRHH electrode was confirmed through its ability to illuminate an LED. The observed luminosity was unaffected even as the FRHH electrode underwent elongation up to 500% of its original length, underscoring its extraordinary ability to maintain conductivity under extensive deformation (Fig. 3c, Video 1†). Samples were subjected to varying degrees of elongation, while monitoring the current at constant voltage. The results indicate minimal variation in current across a strain range of 0 to 18% (Fig. S9†), suggesting that the electrode can accommodate a range of muscle movements including those typically encountered while capturing electrophysiological signals.

For flexible electrodes, a high strain sensing ability is generally undesirable. This is due to the fact that minute skin deformations occurring during movement can exert a substantial influence on the inherent electrical properties of the electrodes, thereby inducing “motion artifacts”. Such artifacts are known to significantly interfere with the accurate extraction and subsequent analysis of the electrical signals.^40–43 Therefore, it is more advantageous for flexible electrodes to exhibit a low strain-sensing ability, coupled with a high electrical conductivity. This combination ensures the high-fidelity transmission of electrophysiological signals, rather than being sensitive to minor skin deformations. To address this concern, we have supplemented Resistance strain curve and GF strain curve of FRHH electrode characterizing the strain response sensitivity of the FRHH electrode (Fig. 3d), the GF value remains below 5 over a 600% strain range. This result indicates that the FRHH electrode exhibits an exceptionally low response to strain stimuli. Therefore, it can be predicted that minor skin deformations will not significantly affect the intrinsic conductivity of the FRHH electrode during practical applications, demonstrating its excellent mechanical stability. The AC impedance of the FRHH electrode after 3000 stretching–releasing cycles (25% strain) was measured to demonstrate the stability of the electrode. The results show that the change in the alternating current impedance values at 1000 Hz is less than 25%, which supports its suitability for long-term wearable applications (Fig. 3e, f and S10†). The uniaxial tensile stress–strain curves of the FRHH electrode under strains ranging from 100% to 600% were measured (Fig. S11†). The results indicate that the FRHH electrode is capable of achieving significant mechanical recovery after the release of tensile forces. The tensile force at 0% and 100% strain for 100 cyclic stretching processes was measured (Fig. 3g). After 100 stretchings at 100% strain, the maximum stress of the FRHH recovers to 95% of its original value (Fig. 3h). The experimental results of the cyclic tensile testing demonstrate that the FRHH electrode exhibits excellent resilience, ensuring reliable mechanical recovery under repeated low-strength deformation. A 4000-cycle bending test (with a bending radius of 1.5 cm) was also conducted, during which the electrode exhibited low resistance changes (ΔR/R₀ < 5%) and did not deviate significantly from its initial value (Fig. S12†). Both the cyclic stretching and bending tests demonstrate the exceptional durability of the electrode.

Fig. S13† visually illustrates the distribution of each component in the FRHH electrode and the conductive pathway schematics under the deformed state. During the stretching process, the two-dimensional nanosheets always overlap with the one-dimensional conductive polymer, forming an electronic conductive pathway. At the same time, the high water content of the hydrogel can continuously release the electrolyte containing Na ions, which can always play the role of ionic conduction. The dual conductive pathway ensures that the FRHH electrode maintains good conductivity during movement, achieving efficient collection of electrophysiological signals. Morphological changes on the surface of the FRHH under different stretching conditions and before and after stretching were observed via SEM (Fig. 3i). The initial state of the hydrogel surface is relatively smooth with uniform porosity. Upon stretching, the texture and micropores became elongated, and no cracking was observed even at a strain of 100%, confirming the alignment and structural integrity of the structure under strain. Importantly, the morphology after recovery is comparable to the initial state, suggesting a high degree of elastic recovery and structural resilience.

A finite element analysis (FEA) of the FRHH electrode was conducted in Abaqus, with the aim of revealing the stress response of the electrode under deformation. Under simulated uniaxial tension ranging from 0 to 30% strain, the stress distribution in the FRHH electrode is primarily concentrated along the SEBS fibers oriented in the x-axis direction, whereas stress in the SEBS fibers oriented along the y-axis direction and in the hybrid hydrogel is extremely low (Fig. S14a and S15†). When uniaxial tension is applied from 0 to 30% along the x-axis, the stress experienced by the electrode is distributed along the x-axis direction in the SEBS fibers, the SEBS fibers in the y-axis direction and stress within the hybrid hydrogel increase linearly with strain. At a strain of 30%, stress in the fibers oriented in the x-axis direction reach 3.3 kPa, which is 1.9 times greater than the stress observed in the hybrid hydrogel. The SEBS fibers in the y-axis direction experience the least amount of stress, amounting to only 1.1 kPa (Fig. S16†). This indicates that during stretching of the FRHH electrode, SEBS fibers oriented along the x-axis direction play a toughening role and alleviate the stress imposed on the hybrid hydrogel by deforming, thereby reducing the likelihood of microcracks appearing in the hybrid hydrogel. Under bending conditions where the minimum bending radius reaches 63.7 mm, the overall force experienced by the FRHH electrode is relatively low (Fig. S14b and S17†). The composite hydrogel at the inner and outer edges is subjected to the highest force, yet the maximum stress is an order of magnitude lower than under tensile conditions. Due to a relatively minor amount of deformation, the stress experienced by the hybrid hydrogel on the neutral plane of the FRHH electrode, as well as the SEBS fibers in both the x-axis and y-axis directions, was lower than on the sides of the hybrid hydrogel (Fig. S18†). This indicates that the FRHH electrode presents a more uniform stress distribution and lower overall stress levels under bending compared to under tensile conditions. The lower stress on the neutral plane and in the SEBS fibers oriented along both axes suggests that the hybrid hydrogel and fiber structure within the FRHH electrodes is well-suited for applications requiring both flexibility and resistance to mechanical stress.

We further performed a series of experiments to demonstrate the practicality of the FRHH electrode in various configurations, such as knotting and twining (Fig. 3j), cutting-reconnection (Fig. 3k), and when submerged in water (Fig. 3l, Video S2†). The luminosity of the LED remained consistent under all conditions, reliably demonstrating the consistent operation of the FRHH electrode with no degradation in functionality. The continuous electrical integrity, notwithstanding various mechanical stresses, demonstrates considerable adaptability. The conductivity stability of the FRHH electrode in both aqueous and saline environments were measured using a four-probe sheet resistance measurement instrument (Fig. S19†). The sheet resistance measurements demonstrate minimal variation even after 120 minutes of immersion, confirming exceptional electrical stability in hydrated physiological conditions. Due to its excellent resilience under a wide range of test conditions, the FRHH electrode is a promising candidate for use in wearable and stretchable electronic devices, particularly where long-term performance is paramount. A comprehensive comparative analysis was conducted to evaluate the AC impedance at 1000 Hz and fracture strain values between the FRHH electrode and various previously reported flexible hydrogel electrodes.^44–53 As quantitatively demonstrated in Table S1,† the FRHH electrode exhibits a substantially reduced AC impedance at 1000 Hz, coupled with significantly enhanced fracture strain values, when compared to existing electrode systems. These comparative data distinctly underscore the superior flexibility and conductivity characteristics inherent to the FRHH electrode architecture.

3.4. Permeability and in vivo biocompatibility tests of FRHH electrode

Biocompatibility is viewed as an essential criterion for wearable devices, however, permeability to water and air is another critical parameter influencing the long-term usability of flexible electrodes, particularly for on-skin applications where breathability is essential. Fig. 4a shows the vapor transmission rate of the FRHH compared against those of conventional flexible electronic substrate materials such as PDMS, Ecoflex and PET films. During a 10-day assessment period, the FRHH electrode demonstrated an enhanced rate of water vapor transmission, demonstrating its exceptional water permeability (Fig. S20†). The water loss amount of hydrogels lacking different single conductive components was measured. The results showed that the difference in weight loss of different experimental groups was very small, which proved that the conductive component had no obvious effect on the water loss of hydrogel (Fig. S21†). The pressure and penetration measurements shown in Fig. 4b, and further confirm the electrode's breathability. Minimal penetration, even under considerable pressure suggests the electrode can form an effective barrier while simultaneously facilitating air flow, thus establishing a skin-compatible interface potentially suitable for extended use. The ion release behavior of the FRHH electrode was analyzed through ICP measurements (Fig. 4c), as indicator of biocompatibility and safety for use in biological applications. The amounts of sodium, sulfur, and titanium ions released after consistent stirring of the FRHH in water for 120 min remained within a biologically tolerable range (Fig. S22†), indicating the electrode may be suitable for prolonged contact with biological tissues. To thoroughly assess the biocompatibility of the FRHH electrode, in vivo tests were conducted on mice (C57BL/6). The FRHH electrode and SEBS fiber control group were directly attached to the skin of mice and secured with medical tape for 2 h, 6 h, or 12 h (Fig. 4d, e and S23†). Histological assessment of the mice skin following exposure to the FRHH electrodes for up to 12 h revealed no discernible inflammatory activity or pathological changes. Collagen fibers within the dermal layers appeared loosely organized and subcutaneous tissue was devoid of any inflammatory cell infiltration (Fig. 4f–h), demonstrating the outstanding biocompatibility of the FRHH electrode.

Fig. 4 Permeability and biocompatibility of the FRHH electrode. (a) Hydrophilicity of the FRHH electrode compared with other commonly used flexible electronic substrate materials; (b) air permeability of the FRHH electrode, Ecoflex film, and PDMS film; (c) concentration of Na⁺, S²⁻, and Ti⁴⁺ ions released by the FRHH electrode after 30, 60, and 120 minutes of underwater stirring; (d) schematic diagram of in vivo experiments used to assess the biocompatibility of the FRHH electrode and SEBS fiber; (e) results of biocompatibility experiment performed on mice skin surface after continuous attachment for 12 h, with no significant changes in local skin before and after attachment; (f–h) immunohistochemical analysis of local skin specimens after 2 h, 6 h, and 12 h of direct contact with the FRHH electrode, showing good tissue condition and no inflammatory response.

3.5. Interface impedance analysis and electrophysiological signals recording

In non-invasive electrophysiological signal acquisition, the electrode-skin contact impedance directly affects signal quality (Fig. 5a).⁵⁴ A dry electrode in contact with the skin constitutes a solid–solid contact without a liquid coupling layer, and their impedance is primarily derived from the stratum corneum resistance (R₂) at the MΩ level, resulting in a high impedance contact. In contrast, when a wet electrode is used, the presence of the electrolyte (or conductive gel) forms a liquid-phase coupling layer, which reduces R₂ by an order of magnitude, bringing it down to the kΩ level, thereby establishing a low impedance contact (Fig. 5b). Therefore, to facilitate the construction of a low-impedance electrode-skin interface, permeating the stratum corneum is of paramount importance. By leveraging the favorable characteristics of hydrogel matrix, the FRHH electrode gradually releases a physiological saline solution (0.9% NaCl) while in contact with the skin. To visually demonstrate the sustained release characteristic of the electrolyte, the surface of the FRHH electrode was placed on a piece of tissue paper was placed and subsequently removed. Upon removal of the electrode, the tissue paper exhibited signs of permeation, indicating the gradual release of the electrolyte from the electrode surface (Fig. 5c). This not only ensures the biosafety of the skin but also effective infiltration of the stratum corneum. To explore the performance of the FRHH when attached to the human body, the impedance of the hydrogel soaked in saline solution was measured (Fig. S24†). It is indicated by the results that the AC impedance at low frequencies is shown to have a slight increase with soaking time, with the peak being approached at 120 minutes, while the high-frequency impedance is remained stable. The stable high-frequency response reflects robust electronic conductivity from the conductive fillers (PEDOT:PSS/Ti₃C₂T_x), ensuring reliable performance during prolonged body contact. To ensure conductivity of the signal acquisition pathway, the impedance between the FRHH electrode and human skin was measured using electrochemical impedance spectral technology (Fig. S25†). The results show a lower AC impedance between the FRHH electrode and human skin compared with those of the dry electrode and commercial wet electrode across a frequency range of 10⁻¹–10⁵ Hz (Fig. 5d, e and S26†). Based on the AC impedance results, we performed circuit modeling of the electrode-skin interfaces using the three types of electrodes. The results revealed that the FRHH electrode system has comparatively lower skin resistance (R_s) and contact resistance (R_c) in contrast to the other two systems (Fig. S27†). This outcome confirms the role of the controlled electrolyte release mechanism in reducing the impedance at the electrode-skin interface, further suggesting that the FRHH electrode is capable of higher performance in sensitive electrophysiological signal recording processes like ECG and EMG.

Fig. 5 Noninvasive electrophysiological signal acquisition using the FRHH electrode. (a) Schematic illustrating the relationship between electrode-skin impedance and skin moisture during EMG and ECG signal acquisition; (b) equivalent circuit diagrams for dry and wet contact between the electrode and skin, with the impedance of each circuit element; (c) demonstration of the slow-release of electrolytes from the FRHH electrode, showing a moistened piece of tissue paper after pressing the FRHH electrode against the tissue and then removing it; (d) impedance spectra of the FRHH electrode, dry electrode, and wet electrode, and (e) their corresponding impedance arcs; (f) ECG signals recorded from a subject at rest using the FRHH electrode, dry electrode, and wet electrode; (g) ECG signals continuously recorded for 20 s during an indoor cycling exercise, (h) a magnified view of the ECG signals at the onset of exercise, highlighting differences in signal-to-noise ratio, motion artifacts, and adhesion among the three types of electrode; (i and j) EMG signals acquired from the FRHH electrode, dry electrode, and wet electrode while placed on a subject's hands, where the FRHH and dry electrodes were placed on the left hand and the FRHH and wet electrodes were placed on the right hand; (k) a four-channel wireless gesture analysis system constructed using FRHH electrodes placed on the subject's forearm targeting four different muscles; (l) four-channel EMG signals collected during different hand gestures, showing a different signal intensity distribution correlates with each gesture due to the muscles involved.

To further evaluate and compare the on-skin electrophysiological signal acquisition capabilities of the electrode, we acquired ready-to-use ECG signals from a subject in a resting state using the FRHH electrode, a dry electrode (control), and a commercial wet electrode. The dry electrode was unable to effectively and continuously record the ECG signal, whereas the FRHH electrode demonstrated a signal quality comparable to that of the commercial wet electrode (Fig. 5f, Video S3†). To assess the stability of the electrode-skin interface under motion conditions, we continuously recorded ECG signals during 20-s intervals from a subject engaged in indoor cycling exercises. The signals obtained from the dry electrode contained motion artifacts and baseline drift across all time periods, primarily due to the solid–solid contact between the electrodes and the skin, which is prone to relative displacement during movement (Fig. 5g, Video S4†). The ECG signal recorded by the FRHH electrode was more stable and had a higher SNR compared to the commercial wet electrode (Fig. 5h). This enhanced quality can be attributed to the matched Young's modulus between the FRHH electrode and the skin, as well as the formation of a more stable electrode-skin contact interface via hydrogen bonding, as mentioned earlier.

EMG signals are another important class of electrophysiological signals closely related to human movement. To control for variables, we constructed a four-channel system on both forearms of a single human subject to synchronously record EMG signals during continuous fist clenching (Fig. 5i). Signals were acquired using the FRHH electrode, dry electrode, and commercial wet electrode. The FRHH electrodes consistently captured the highest quality EMG signals from both the left and right forearms. In contrast, the dry electrode experienced a continuous decline in SNR after the subject's initial clench due to contact issues. By the sixth clench, the EMG information had become indistinguishable (Fig. 5j). A video of the simultaneous recording of the EMG signals captured by the FRHH electrode and commercial electrode on different arms was recorded (Video S5†). The analysis clearly reveals that the SNR of the signals captured by the FRHH electrodes is superior to that of the commercial electrode. In addition, control experiments were conducted using pure PVA hydrogel without conductive fillers. As shown in Fig. S28,† the pure PVA hydrogel exhibited significantly lower SNR compared to the FRHH electrode (with PEDOT:PSS/Ti₃C₂T_x), The results indicate that in the absence of PEDOT:PSS/Ti₃C₂T_x addition, the PVA hydrogel solely possesses ionic conductivity, impeding the efficient transmission of electrophysiological signals. Once PEDOT:PSS/Ti₃C₂T_x is added, electron-conducting channels are introduced, thus enhancing the efficiency of electrical signal transmission. Finger movements are primarily driven and controlled by the forearm muscle groups, therefore, effectively acquiring EMG signals from the corresponding forearm muscles not only holds promise for the diagnosis and treatment of finger movement disorders, but also for precise gesture recognition. The FRHH electrode was used to construct a four-channel wireless EMG recording system on the forearm, targeting muscles including the flexor digitorum superficialis, flexor carpi radialis, cubitalis gracilis, and flexor carpi ulnaris, to test the EMG-based gesture recognition (Fig. 5k). Significant differences in the intensity of the four-channel EMG signals were observed when subjects performed various hand gestures. The primary reason for these variations is the selective engagement and differing levels of muscle activation involved in different gestures (Fig. 5l). Compared to gesture recognition systems based on traditional strain or stress sensors, an EMG-signal-based gesture recognition system can precisely reflect a subject's intent solely based on electrophysiological muscle activation information provided during the finger movement process. This approach holds potential value for medical research and rehabilitative training applications.

4. Conclusion

In summary, we have reported a FRHH electrode with exceptional mechanical and electrical properties, making it suitable for high-quality on-skin electrophysiological signal recording. The FRHH electrode combines the attributes of electrospun SEBS fibers, PEDOT:PSS, Ti₃C₂T_x, and a PVA hydrogel matrix to achieve a resulting substrate with excellent stretchability, high conductivity, and good biocompatibility. The controlled release of electrolytes can significantly reduce impedance at the electrode-skin interface, thereby enhancing the SNR of the acquired electrophysiological signals. Our experimental results demonstrated stable and accurate recording of ECG and EMG signals on a human subject, even under dynamic conditions. Furthermore, the FRHH electrode shows great potential in gesture recognition, with the potential to provide a more intuitive and precise method for interpreting muscle activation patterns. The biocompatibility and non-irritant attributes of the FRHH electrode were demonstrated through rigorous in vivo evaluations, and the proposed electrode is therefore a viable contender for prolonged use in wearable electronics applications such as health status monitoring and analysis and exercise monitoring.

Data availability

The data supporting this article have been included as part of the ESI.†

Conflicts of interest

There is no conflict of interest to declare.

Acknowledgements

This work was financially supported by the National Natural Science Foundations of China (No. 32127801, 62104051), and Natural Science Foundation of Guangxi Province (No. AE31200115).

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Footnote

† Electronic supplementary information (ESI) available. See DOI: https://doi.org/10.1039/d4ta08229j

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