Tailored construction and functional applications of conductive hydrogels for bioelectronic interfaces

Abstract

Conductive hydrogels have emerged as core foundational materials for bioelectronic interfaces, uniquely bridging the mechanical and electrical property gap between biological tissues and electronic devices. Unlike previous reviews that primarily focus on material classification and broad device applications, this review adopts an interface-centric perspective, specifically targeting the performance optimization of conductive hydrogels for bioelectronic interface applications. We first categorize conductive hydrogels according to their underlying conductive mechanisms, and describe the morphology and interfacial interactions of conductive nanomaterials with different dimensionalities. We then systematically examine interface-relevant design strategies: (i) achieving conformal contact by reducing stiffness, enhancing adhesion, and decreasing thickness; (ii) constructing a reliable and stable interface by improving toughness and imparting self-healing properties; and (iii) enhancing signal transmission efficiency through the synergistic combination of high conductivity, efficient charge transfer, and low interfacial impedance. We further review recent advances in flexible bioelectronic devices, and finally discuss current challenges and future directions, including clinical translation, device miniaturization, and integration into fully automated bioelectronic systems.

Introduction

Over the past decade, bioelectronic devices featuring flexibility, stretchability and high reliability have achieved rapid development. Flexible devices feature tunable modulus from kPa to MPa [1–3], achieve high stretchability and have favorable bending adaptability [4–6]. These devices effectively overcome the limitations of rigid materials, including fixed morphology, brittleness and failure to adapt to complex curved surfaces [7, 8]. They can conformally attach to complex surfaces, including human skin and internal tissues. This greatly reduces mechanical mismatch at bio-abiotic interfaces. They provide diverse flexible sensing solutions for numerous application fields, including healthcare [9, 10], flexible displays [11], and wearable devices [12, 13].

Currently, bioelectronic devices integrate flexible materials into electronic systems through diverse assembly strategies to meet specific functional demands. These flexible materials possess biocompatibility, mechanical compliance, and stability, with typical examples including conductive hydrogels, metal films, conductive elastomer composites, and conductive textiles [14, 15]. Among these, conductive hydrogels, as a new class of flexible conductive materials, exhibit unique advantages. They are highly hydrated 3D networks formed by physical or chemical crosslinking of polymers. In nature, electrical conductivity originates from the transport of electrons or ions. Therefore, the construction of conductive gels is based on introducing electronic conductors and ions into the gel matrix [16, 17]. Conductive media (e.g., ions, conductive polymers, carbon nanomaterials, and metal nanostructures) are incorporated into these networks, thereby endowing the hydrogels with conductivity. Conductive hydrogels combine the softness and the charge transport capability of conductive materials.

Compared with conventional flexible electronic materials such as metal films, conductive elastomer composites and conductive textiles, conductive hydrogels show comprehensive advantages in flexible bioelectronic devices owing to their modulus match, biocompatibility and environmental stability. First, conductive hydrogels possess a tunable Young’s modulus in the kPa range [18], enabling conformal contact with soft biological tissues including skin, nerves and muscles. Second, the highly hydrated 3D network structure confers biocompatibility and tissue affinity, rendering them ideal for long-term biointerface applications. In addition, compositional tuning can endow hydrogels with antidehydration and antifreezing properties, enabling them to maintain stable electrical and mechanical performance in dry and low-temperature environments [19–21]. Collectively, conductive hydrogels achieve synergistic optimization of mechanical compliance, biocompatibility, stability, and conductive reliability under large deformation. They have emerged as one of the most promising flexible conductive material systems for wearable health monitoring, implantable neural electrodes, human-machine interfaces, flexible sensors, and other related fields [22–25].

In recent years, significant progress has been achieved in the design strategies of conductive hydrogels [26]. Based on a systematic literature search in Web of Science, Google Scholar, and PubMed using keywords such as “conductive hydrogel,” “bioelectronic interface,” and “interfacial impedance,” this review focuses on the performance optimization of conductive hydrogels for bioelectronic interface applications. Recent reviews have primarily addressed the classification of conductive hydrogels, their multifunctional properties, and broad device applications. For instance, He et al. classified conductive hydrogels into ionic and electronic conductive types and highlighted their applications in wearable devices, tissue engineering, and bionic robots [27]. Zhu et al. reviewed multifunctional conductive hydrogels, emphasizing functionalities such as self-healing, adhesion, anti-swelling, and stimuli responsiveness, and illustrated their use in strain sensors, supercapacitors, touch panels, and bioelectronic devices [17]. Another review by Zhu et al. summarized synthesis strategies for conductive hydrogels and their applications in flexible neural electrodes, sensors, energy storage devices, and soft robots [28]. Mo et al. specifically reviewed conductive hydrogel-based wearable sensors, focusing on sensing performance and healthcare monitoring [29]. These reviews generally treat bioelectronic interfaces as one of many application scenarios rather than as a central design target. Consequently, targeted discussions on the core performance requirements of bioelectronic interfaces, such as conformal contact, ion–electron transduction, interfacial impedance, and long-term electrochemical stability, remain insufficient.

Against this background, this review specifically focuses on the performance optimization of conductive hydrogels for bioelectronic interface applications. We first categorize conductive hydrogels according to their underlying conductive mechanisms, and further describe the nanoscale dimensions, morphology and interfacial interaction of conductive nanomaterials with different dimensionalities. We then systematically examine interface-relevant design strategies: (i) achieving conformal contact by reducing stiffness, enhancing adhesion, and decreasing thickness; (ii) constructing a reliable and stable interface by improving toughness and imparting self-healing properties; and (iii) enhancing signal transmission efficiency through the synergistic combination of high conductivity, efficient charge transfer, and low interfacial impedance. We further review recent advances in flexible bioelectronic devices, and finally discuss current bottlenecks and future research prospects in depth.

Classification of conductive hydrogels according to conductive mechanism

Depending on the underlying conductive mechanisms, CHs can be categorized into ionic CHs (ICHs), electronic CHs (ECHs), and hybrid CHs [27].

ICHs

ICHs are conductive hydrogels that use inorganic electrolyte salts, polymerizable ionic monomers, zwitterionic polymers, etc., as ion sources (Fig. 1a-c). Ionic hydrogels are cross-linked polymer networks capable of absorbing and retaining large amounts of water or biological fluids, containing significant concentrations of charged groups (ions) within their structure. ICHs rely on ion‑transport channels and aqueous environments to facilitate efficient ion conduction, and this process can be further enhanced by optimizing pore distribution and increasing electrolyte concentration. The ionic conductors in ICHs can generally be divided into three categories: electrolytes (Li⁺, Na⁺, Ca²⁺, Al³⁺, Fe³⁺ and K⁺), polyelectrolytes, and ionic liquids(ILs) (e.g., 1‑ethyl‑3‑methylimidazolium ([EMIM]⁺) and 1‑vinyl‑3‑butylimidazolium ([VBIM]⁺)). Among various ion‑conductive hydrogels, electrolyte‑based ICHs are a common and representative type [30–32]. An electrolyte is generally defined as a substance that dissociates into ions when dissolved in a solution. For instance, researchers have developed a sodium salt-based ICH, which exhibits comprehensive mechanical properties: an ultrahigh stretchability of 2167%, a favorable tensile strength of 19.2 kPa, and a conductivity of 1.71 S·m^-1 [32]. Different from electrolyte‑based ICHs that are usually doped with inorganic metal salts, polyelectrolyte‑based ICHs represent another type of materials. The polyelectrolyte is an organic compound characterized by charged side groups attached to its polymer chains. Commonly used polyelectrolytes in hydrogels include PAA, polystyrene sulfonate, and natural polyelectrolytes such as chitosan. Researchers fabricated a poly(acrylic acid)/chitosan double-network hydrogel. [All-in-one” chitosan modified polymerizable water-in-eutectic gel for human signal detection in various scenarios] ILs are also frequently incorporated into hydrogels as a distinctive type of electrolyte to enhance their electrical conductivity and other functional properties [33,34]. Composed of organic cations and anions, ILs generally exist as molten organic salts with melting temperatures below 100 °C. These liquids exhibit a range of intriguing physicchemical properties, including ionic conductivity, thermal stability, non‑flammability, and a broad electrochemical stability window. Numerous different ILs can be synthesized from a vast pool of potential cation‑anion combinations, and each IL possesses its own set of functions. Researchers have developed a highly stable ionic hydrogel. ILs enhance the environmental adaptability of the hydrogel, with key properties as follows: water retention rate > 90% after 7 days, freezing tolerance below − 70 °C, and ionic conductivity up to 1.2 S·m⁻¹ [35].

Fig. 1.

Classification of conductive hydrogels according to conductive mechanism. a-c ICHs (Ions, ILs, Polyelectrolyte); d-f ECHs(conductive polymers(PEDOT, PANI PPy)); g-i ECHs(additive nanomaterials): 0D(nanoparticles), 1D(CNTs), 2D(rGO)

ECHs

Compared with ICHs, ECHs conduct electricity via electron transfer. By incorporating conductive fillers (such as metals and carbon‑based nanomaterials) or conductive polymers into cross‑linked hydrogel networks, ECHs possess high electrical conductivity and biomechanical compatibility. This electron‑based conduction enables faster and more efficient charge transfer, as it circumvents the inherent limitations of ion migration, such as viscosity‑dependent ion motion and incomplete ion dissociation [36]. We further classify them into two types: conductive polymer hydrogels (Fig. 1d-f) and conductive hydrogel with additive nanomaterials (Fig. 1g-i).

Conductive polymer hydrogels

Conductive polymers (CPs) directly construct conductive hydrogels through their conjugated polymer chains with alternating single‑ and double‑bond configurations [37]. The three‑dimensionally interconnected framework within hydrogels builds uninterrupted conductive pathways to enable fast electron transport.

Classic representatives of such electron‑conductive polymers include three conjugated systems: poly(3,4‑ethylenedioxythiophene) (PEDOT), polypyrrole (PPy), and polyaniline (PANI). PEDOT possesses a highly planar conjugated backbone, delivering exceptional intrinsic electrical conductivity and long‑term environmental stability. It is typically compounded with poly(styrene sulfonate) (PSS) to further improve conductivity and processability, and the PEDOT: PSS system will be systematically elaborated in the subsequent Sect. [38]. PPy is a typical intrinsic conductive polymer with extensively π‑conjugated molecular chains, characterized by tunable conductivity as well as favorable thermal and environmental stability [39]. PANI is distinguished by its facile synthesis route, low‑cost monomers, and regulated doping/dedoping behavior, which allows flexible adjustment of its conductivity over a wide range [40]. Nevertheless, conductive hydrogels formed via direct cross-linking of pristine conductive polymers generally boast high intrinsic conductivity but suffer from inferior mechanical performance [17]. To address this shortcoming, diverse fabrication strategies have been developed to engineer conductive hydrogels that combine high electrical conductivity with excellent stretchability. For example, researchers employed S-EDOT bearing sulfonic acid side groups, and the resulting hydrogel achieved a high toughness of 620 kJ·m⁻³ and a high electrical conductivity of 580 S·m⁻¹ [41].

Conductive hydrogels with additive nanomaterials

Conductive hydrogels formed by direct cross-linking of conductive polymers usually exhibit high conductivity but possess poor mechanical properties [17]. Incorporating conductive nanofillers into hydrogel networks is currently a core strategy for synergistically optimizing the electrical and mechanical properties of hydrogels [27]. Such conductive fillers cover various systems including metals, carbon‑based materials, and transition‑metal carbides/nitrides, and can be categorized by spatial dimension into three types: zero‑dimensional nanoparticles, one‑dimensional nanowires, and two‑dimensional nanosheets. The dimension, micromorphology, interfacial interactions, and size‑dependent effects of nanomaterials significantly influence the electrical performance, percolation threshold and other properties of composite systems [42].

(1) Conductive hydrogels with additive 0D nanomaterials

0D nanomaterials are a widely used class of conductive fillers in conductive hydrogels, covering various material systems such as metallic elements (gold, silver) [43, 44] and metal oxides (alumina, cerium dioxide, zinc oxide, tin oxide). They share the common feature that all three‑dimensional spatial dimensions are at the nanoscale. They can be incorporated into hydrogel systems through multiple approaches including solution blending and in‑situ reduction [45], dispersing and filling the gaps of polymer networks with little disturbance to the hydrogel network structure. Nevertheless, their granular morphology brings about point‑to‑point contact between particles, leading to high interfacial charge‑transfer resistance and higher filler loading required for constructing continuous electron pathways. Hence, their percolation threshold is higher than that of one‑dimensional and two‑dimensional filler systems [46]. In addition, the interfacial bonding strength between particles and the matrix directly governs stress‑transfer efficiency. Unmodified metal nanoparticles bind to polymer chains merely via weak van der Waals forces, so interfacial debonding readily occurs under external force, resulting in very limited mechanical reinforcement. Enhancing interfacial bonding via in‑situ reactions or surface modification can simultaneously improve interfacial charge‑transfer efficiency and mechanical load‑bearing capacity. For instance, in the lignin‑based supramolecular hydrogel fabricated, silver nanoparticles were in‑situ generated from silver ions through dynamic redox reactions with PEG‑LGS. Tight interfacial bonding was formed between nanoparticles and polymer chains. Combined with host‑guest interactions between α‑cyclodextrin and PEG‑LGS to enrich cross‑linking sites, the hydrogel finally achieved an adhesion strength of 48 kPa and a fracture strain increased to 1390%. Furthermore, the hydrogel exhibited a high electrical conductivity of up to 2.71 S m − 1 and demonstrated excellent self-healing capabilities [47]. In another study, researcher used silver nanoparticle (AgNP) solution instead of pure water as the reaction solvent, and then immersed POAPZ, a polyvinyl alcohol/oxidized carboxymethyl chitosan (PVA/O‑CMCS) -based composite hydrogel, into the mixed solution of PDO‑zinc sulfate (ZnSO₄). Under the synergistic effects of cross‑linking coordination, salting‑out and hydrogen bonding, the finally prepared hydrogel exhibited excellent electrical conductivity (with an electrical conductivity of 1.4 S·m⁻¹) and mechanical properties, and its fracture stress can reach 2.2 MPa [48].

(2) Conductive hydrogels with additive 1D nanomaterials

1D nanofillers refer to linear conductive materials whose two radial dimensions are on the nanoscale while their axial dimension can extend to micrometers or even millimeters. They cover typical systems such as metal nanowires (e.g., silver nanowires, AgNWs) [49] and carbon nanotubes (CNTs), and can be incorporated into hydrogel systems via various approaches including solution blending, in‑situ growth, and vacuum‑assisted impregnation [50]. Unlike 0D nanomaterials, the linear morphology endows fillers with line‑to‑line overlapping contact for electrical conduction. The effective contact area between fillers increases and the interfacial charge‑transfer resistance decreases. The filler content required to form continuous electronic pathways is lower than that for zero‑dimensional particle systems. Hence their percolation threshold is lower than that of zero‑dimensional filler systems. Among them, CNTs and silver nanowires, with high aspect ratios and intrinsic flexibility, tend to entangle with one another and are capable of constructing continuous percolative conductive networks with higher interconnectivity [51, 52]. Apart from this, the interfacial bonding strength between fillers and the matrix also directly affects the stress transfer efficiency. One‑dimensional nanostructures with a high aspect ratio can dissipate fracture energy through mechanisms such as crack deflection, effectively transfer the stress borne by the matrix, and improve both the strength and toughness of hydrogels [53]. Nevertheless, these 1D nanofillers have limitations. For instance, they exhibit poor dispersibility and tend to agglomerate within the hydrogel matrix, giving rise to localized stress concentration and consequently rendering the hydrogel brittle. Furthermore, considerable contact resistance exists among nanowires, which usually limits the conductivity and stability of conductive hydrogels based on silver 1D nanofillers. If surface functional modification or in‑situ growth strategies are adopted to strengthen the interfacial interaction between fillers and the polymer matrix while improving the dispersion uniformity of fillers, the interfacial charge‑transfer efficiency and mechanical load‑bearing capacity can be optimized simultaneously. For instance, by modifying the silver nanowires, the dispersibility and conductivity in polyvinyl alcohol were improved [54].

(3) Conductive hydrogels with additive 2D nanomaterials

Owing to their unique electronic structures and ultra‑high specific surface areas, 2D nanomaterials have emerged as cutting‑edge material systems with great application potential in numerous fields. Starting from graphene, the first‑discovered material of this kind, to subsequently developed systems such as transition‑metal dichalcogenides, phosphorene, and MXene, these materials exhibit outstanding properties including excellent electrical conductivity and mechanical strength. Graphene possesses a two‑dimensional hexagonal lattice molecular structure similar to that of carbon nanotubes, and shows superior intrinsic electrical conductivity and mechanical properties. Nevertheless, its intrinsic hydrophobicity prevents uniform dispersion in aqueous phases, which restricts its direct application in conductive hydrogels. Graphene oxide (GO), the most widely used graphene derivative, carries abundant hydrophilic functional groups such as hydroxyl and carbonyl groups on its sheet surface. These groups greatly strengthen the interactions between sheets, water molecules and polymer chains, endowing GO with favourable water dispersibility and interfacial compatibility. However, oxygen‑containing groups disrupt the intrinsic sp² conjugated structure of graphene, hinder inter‑sheet electron transport, and thus cause a substantial drop in electrical conductivity [55, 56]. Reduced graphene oxide (rGO) can be obtained by reducing GO via processes such as high‑temperature treatment and dopamine reduction. The reduction process restores the sp² conductive structure to a certain extent, enabling the material to achieve excellent intrinsic electrical conductivity. Meanwhile, an appropriate amount of oxygen‑containing functional groups can be retained on sheet surfaces to maintain stable aqueous dispersibility. With a large specific surface area and desirable mechanical properties, rGO serves as the most mainstream filler for graphene‑based conductive hydrogels [57, 58]. As a two‑dimensional filler, rGO can form a complete percolative conductive network at relatively low loading content. In addition, its high‑aspect‑ratio sheet structure dissipates fracture energy through mechanisms including inter‑layer slipping, crack deflection and sheet pull‑out, and therefore holds significant potential for enhancing the strength and toughness of hydrogels. Even so, rGO fillers have inherent limitations [59]. Strong van der Waals forces and π‑π interactions between sheets readily induce stacking and agglomeration, making uniform single‑sheet dispersion hard to realise within hydrogel matrices and impairing the continuity of conductive networks. Besides, residual oxygen‑containing groups on rGO surfaces still partially damage conductive structures, so rGO can hardly reach the conductivity level of intrinsic graphene. Furthermore, unmodified rGO, like carbon‑based materials of other dimensions, has weak interfacial bonding force with polymer chains. Surface functional modification can improve sheet dispersibility and reinforce interfacial interactions between fillers and polymer matrices, so as to optimise both the electrical and mechanical performances of composite hydrogels. For instance, inspired by the biomimetic modification strategy of dopamine, researchers fabricated polydopamine‑coated rGO reinforced composite conductive hydrogels through thermal polymerization. Polydopamine modification not only improved the dispersibility and interfacial compatibility of rGO, but also constructed dynamic covalent and non‑covalent cross‑linking networks within hydrogel matrices, which consist of C‑N covalent bonds, intra‑/inter‑molecular hydrogen bonds, NH₃⁺‑π interactions and π‑π interactions. The synergy of multiple interfacial interactions endows the hydrogel with excellent comprehensive performances, featuring a fracture stress of 496.4 ± 85.9 kPa, an elongation at break of 1156.9 ± 110.1%, as well as favorable electrical conductivity [60].

MXenes are a new class of 2D early transition‑metal carbon/nitride materials fabricated by etching and exfoliating layered MAX‑phase precursors. They possesses multiple advantages including large specific surface area, high intrinsic electrical conductivity, inherent hydrophilicity and flexibility [61–63]. Compared with graphene‑based fillers, MXene sheets are naturally decorated with abundant terminal functional groups such as ‑OH, ‑O and ‑F. Strong interfacial interactions can be formed with hydrogel polymer matrices without additional modification. Meanwhile, their intrinsic metallic‑level conductivity endows composite hydrogels with favorable charge‑transport capability. As 2D conductive fillers, MXene nanosheets feature a low percolation threshold and high conductive efficiency. Nevertheless, MXene suffers from inevitable oxidative degradation in aqueous media, which gives rise to gradual deterioration of conductive performance and severely limits the long‑term service stability of hydrogels. Existing research strategies can mitigate oxidation by surface modification with hydrophobic groups to reduce contact sites between nanosheets and water molecules. For instance, to address the oxidation‑stability issue of MXene, researcher designed composite hydrogels based on catechol‑functionalized poly(vinyl alcohol) (PVA‑CA). Hydrophobic catechol groups can bind sufficiently to the MXene surface and reduce water‑reactive oxidation sites on the sheets, thereby effectively suppressing the oxidative degradation of MXene inside hydrogels. Furthermore, strong interfacial hydrogen bonds are formed between MXene nanosheets and PVA‑CA polymer chains. Combined with the cross‑linking effect between borate anions and polymer hydroxyl groups, the hydrogel acquires favorable self‑healing capacity, enabling rapid self‑repair and performance recovery after mechanical damage [64].

Hybrid conductive hydrogels

Hybrid conductive hydrogels couple ionic and electronic conduction pathways to achieve enhanced electrical conductivity, superior mechanical properties, and expanded functional versatility [65]. Among the widely explored hybrid CH systems, composites of conductive polymers and acidic dopants represent a major category, where the doping mechanism serves as the core driver for conductivity enhancement. A paradigmatic example is the PEDOT: PSS hydrogel prepared by a surface gelation method coupled with chemical acid treatment and controlled partial densification, which deliver an ultrahigh conductivity of up to 1,200 S cm⁻¹ and a tensile strength reaching 30 MPa [66]. Specifically, acid treatment exerts multiple reinforcing effects: it protonates PSS polyelectrolyte chains, eliminates metal-ion crosslinkers, and strengthens interchain π–π stacking and hydrophobic interactions between PEDOT segments. The follow-up controlled densification further tailors the hydrogel microstructure by collapsing micropores into nanopores, optimizing pore architecture and promoting electric double-layer formation, thus enabling independent tuning of mechanical and electrical performance[].

In summary, ICHs exhibit excellent biocompatibility and adaptability for soft electronic applications, while ECHs feature superior charge transport efficiency. Hybrid CHs integrate the merits of both conduction modes, offering a practical approach to constructing multifunctional hydrogels with enhanced comprehensive performance.

Core engineering strategies for optimizing conductive hydrogel-biological tissue interfaces

Conformal contact

Achieving conformal contact is one of the core indicators determining the application performance of conductive hydrogels in flexible electronics, biomedical sensing, and other related fields [67]. Conformal contact ensures efficient signal transmission and reduces interface interference. It also alleviates biological tissue irritation and improves device wearing comfort. Conformal contact can avoid problems such as signal attenuation and response delay caused by interface gaps. However, most conductive hydrogels still have common problems such as poor contact with the interface, inadequate mechanical compliance, and easy detachment. To solve the above problems, researchers have designed a variety of construction strategies [68–71]. Currently, the mainstream strategies for achieving seamless interfacial contact mainly include modulus matching with biological tissues, ultrathin structural design, and interfacial adhesion optimization. These approaches enable precise regulation of hydrogel modulus, film thickness, and surface chemical properties, thereby strengthening interfacial interactions between hydrogels and biological tissues. Such synergistic modulation facilitates robust conformal contact, which in turn significantly improves electrical signal transmission efficiency and long‑term operational stability [68, 70].

Modulating modulus is a fundamental strategy to achieve conformal contact. The underlying mechanism involves tailoring hydrogel crosslinking density, incorporating flexible segments, or constructing dynamically crosslinked networks to modulate hydrogel modulus and achieve better conformity with target tissues [72–75]. For example, incorporating flexible polymer segments (e.g., polyethylene glycol (PEG) and polypropylene glycol (PPG)) can effectively lower hydrogel stiffness, enhance material flexibility and mechanical compliance, and alleviate interfacial stress concentration caused by modulus mismatch. This can help reduce interfacial gaps during contact and reduce mechanical irritation to biological tissues, thereby facilitating conformal contact between hydrogels and target interfaces and providing a structural basis for stable biointerfacial integration. Specifically, tannic acid modification tailors the 3D network topology of aramid nanofiber (ANF)–polyvinyl alcohol (PVA) composites to construct an ultrathin hydrogel film. While maintaining high mechanical strength, this film achieves a modulus of 600 kPa which is comparable to the range reported for natural human skin, enabling tunable compliance and conformal adaptation to biological organs [76] (Fig. 2a-b).

Fig. 2.

Strategies for achieving conformal contact.a Schematic illustration of the tannic acid-regulated 3D network topology of ANF-PVA; b Schematic diagram of the electrode-integrated structure of the ultrathin hydrogel film; c Cross-sectional scanning electron microscopy (SEM) image of the hydrogel nanofilm. Scale bar: 2 μm; d SEM image of the organohydrogel nanofilm conforming to a fingerprint replica, illustrating its ultrathin structure and high compliance. Scale bar, 500 μm. The inset displays an optical microscope image revealing its nanomesh-reinforced architecture. Scale bar, 5 μm; e Microscopic image showing the conformal conformal contact between the nanofilm and the surface of human skin. Scale bar: 500 μm; f Adhesion of device on skin; g Optical images of bioadhesive hydrogel; h Adhesion force of the hydrogel according to glycerol loading change (n = 5, independent samples)

Reducing thickness represents another effective strategy for optimizing conformal interfacial contact. This strategy works by tailoring fabrication processes to minimize hydrogel thickness, which in turn reduces thickness-induced interfacial gaps and strengthens interfacial conformity [77–79]. Ultrathin conductive hydrogel films with tunable thickness can be readily fabricated through solution casting, electrospinning, layer-by-layer self-assembly, and template-assisted methods [78, 80–82]. These ultrathin materials conform to the microtopography of the contact interface and eliminate interfacial gaps at both macro- and microscales. They also reduce the overall device dimensions, making them ideal for wearable and implantable applications with strict size constraints. Furthermore, ultrathin design alleviates problems such as insufficient conformal contact and easy delamination caused by excessive thickness, thereby further boosting seamless interfacial performance [76]. For example, the 392-nanometer-thick hydrogel fabricated via stepwise dip-coating of a polyurethane nanofibrous network can mimic skin deformation and provide a robust substrate platform for skin-integrated bioelectronic devices [83] (Fig. 2c-e).

Modulating surface chemical properties represents another key strategy to maintain stable conformal interfacial contact. Although ultrathin devices may achieve initial tissue adhesion through non-covalent physical interactions, such adhesion is often insufficient for precise positioning and long-term reliable monitoring under physiological conditions. Adhesion is a supramolecular phenomenon in which two adherends are interconnected via covalent bonds, non-covalent interactions, polymer chains, networks, or nanoparticles [84, 85]. During detachment, energy dissipation occurs through multiscale cascading processes, including bond cleavage, chain retraction, and bulk dissipation. Current strategies for enhancing interfacial bonding strength between hydrogels and contact interfaces include grafting adhesive moieties onto the hydrogel surface, introducing dynamic interfacial adhesion, and forming interfacial covalent linkages. For example, grafting highly adhesive substances such as dopamine, chitosan, and polydopamine on the hydrogel surface can form a firm combination with the contact interface through hydrogen bonds, hydrophobic interactions, and electrostatic interactions. Alternatively, reactions such as photoinitiation and enzymatic crosslinking can be introduced to form covalent bonds between hydrogels and the contact interface. This strategy significantly enhances adhesion stability and resists external disturbances including friction and deformation. It prevents separation between hydrogels and the contact interface. Accordingly, a conformal contact state can be maintained over a long period, ensuring the stable operation of the device [86, 87] (Fig. 2f-h). As a representative example, catechol-conjugated alginate (Alg-CA) has been used to synthesize tissue-adhesive hydrogels. Serving as the matrix for flexible devices, such hydrogels can firmly attach to biological tissues and deliver reliable biosignal monitoring performance [88]. They achieve strong tissue adhesion (1–3 kPa), which favors conformal contact and supports accurate monitoring both in vitro and in vivo.

Stability and reliability

While conformal contact ensures efficient interfacial coupling, long-term service stability and mechanical reliability constitute another prerequisite for the practical application of conductive hydrogels [89]. In scenarios such as wearable sensing and implantable devices, conductive hydrogels need to withstand long-term external mechanical disturbances, environmental changes (e.g., humidity and temperature fluctuations), and cyclic service wear. Stable structural and conductive properties ensure consistent signal transmission, while reliable performance prevents device failure caused by material degradation. Currently, most conductive hydrogels are still plagued by inherent drawbacks, including inadequate toughness, susceptibility to fracture, and poor self-recovery capability after mechanical damage. These issues collectively lead to degraded structural and functional stability, as well as insufficient long‑term operational reliability. Therefore, developing conductive hydrogels with outstanding stability and reliability is essential to promote their industrial translation. To overcome these limitations, researchers have proposed diverse material design strategies. Among them, enhancing mechanical toughness and endowing hydrogels with self‑healing ability can equip the materials with robust structural support and autonomous damage‑repair capabilities [53, 90]. By optimizing the energy dissipation efficiency and dynamic responsive features of the gel network, these modifications strengthen structural stability, suppress damage initiation, and enable autonomous self‑repair of damaged regions. Such synergistic regulation further improves the stability and service reliability of conductive hydrogels in practical scenarios.

Enhancing toughness is a fundamental strategy to strengthen the stability and reliability of conductive hydrogels [91]. The underlying mechanism involves the construction of an efficient energy-dissipating network [92, 93]. Typical toughening methods include the construction of double-network structures [94–97], mechanical training [98, 99], crystallization regulation [30], and solid-phase doping [100, 101]. The double-network structures introduce sacrificial bonds. Mechanical stretching can induce the orientation of polymer chains, which effectively suppresses crack propagation and improves the toughness of hydrogels. The formation of crystallization and solid-phase doping can suppress crack propagation because they have high chain density, and the energy for rupturing a crystalline domain is higher than that for rupturing the amorphous polymers. Meanwhile, high entanglement also serves as a novel strategy to enhance the toughness of hydrogels (Fig. 3a). These strategies enhance the resistance of gel to mechanical damage, suppress structural fracture induced by cyclic utilization and external perturbations, and achieve long-term structural and functional stability. In recent investigations, researchers developed a strategy to address the trade-off between electrical conductivity and stretchability in conductive polymer hydrogels. By modulating the mass ratio of PEDOT: PSS to PVA coupled with acetic acid-assisted treatment, a PEDOT: PSS/PVA double-network hydrogel was fabricated [102]. This hydrogel demonstrated favorable electrical conductivity, exceptional fracture strain (up to ~ 150%), high biocompatibility, suppressed swelling ratio, and outstanding electrochemical properties.

Fig. 3.

Strategies for achieving stability and reliability. a Schematic illustration of the preparation method of the highly entangled double network (HEDN) hydrogels; b Schematic diagram illustrating the ion complexation and crystallization-induced construction strategy for Cry-gel; c Chemical structures of ANF and PVA and their intermolecular hydrogen bonding interactions; schematic illustration of the fabrication process of ACH, including stretching and confined drying processes for regulating the orientation of nanofiber assemblies

At the present stage, researchers have developed a variety of hydrogel strategies with high fracture toughness, such as those regulated by crystalline domains. Inspired by the energy dissipation mechanism of spider silk, researchers designed a double-network conductive hydrogel [103]. The design promoted polymer chain entanglement and the formation of PVA nanocrystalline domains, while inducing horizontal orientation of MXene to form a dense layered structure. Subsequently, competitive destruction of hydrogen bonds within PVA crystalline domains reduced the crystallite size. This process yielded a three-level hierarchical network composed of a micron-scale layered framework, nanometer-scale refined crystalline domains, and molecular-scale dynamic hydrogen bonds. This hydrogel achieved ultrahigh strength (13.4 MPa), ultimate fracture strain (2337%), and toughness (164.6 MJ·m⁻³). In another study, researchers proposed a synergistic strategy that combines dual-ion complexation and crystallization-driven molecular assembly to fabricate a cellulose-based ionic hydrogel. This gel exhibits a maximum tensile stress of 2.3 MPa, an elongation at break of 150%, and an ionic conductivity of 96.8 mS·cm⁻¹ at room temperature [104] (Fig. 3b). Furthermore, a highly anisotropic, multifunctional tendon-mimetic hydrogel was fabricated by leveraging the interfacial interactions between rigid aramid nanofibers and flexible PVA [105]. The hard nanofibers and soft PVA in these anisotropic composite hydrogels simulate the structural interactions between aligned collagen fibers and proteoglycans in natural tendons. This unique biomimetic structure endows the hydrogel with exceptional mechanical properties: a high modulus of 1.1 GPa, a tensile strength of 72 MPa, and a fracture toughness of 7333 J·m⁻² (Fig. 3c).

Although improving material toughness can significantly enhance reliability, there is still a high risk of fracture under repeated stretching and high-strain environments. Therefore, researchers have proposed the strategy of imparting self-healing properties to hydrogels. Self‑healing hydrogels are all based on a common principle — the so‑called “mobile‑phase” Beyond bulk elemechanism, which achieves crack closure by combining mass transfer with the reconnection of broken chain segments within the hydrogel matrix [106] (Fig. 4a). Reconnection within such matrices is usually mediated by non‑covalent or covalent bonds. Non‑covalent interactions rely on weak sacrificial linkages such as ionic bonds, hydrogen bonds [107, 108] or hydrophobic bonds(Fig. 4b), whereas covalent bonds depend on dynamic covalent bonds and metal coordination bonds [109] (Fig. 4c). The self-healing gel can autonomously repair damaged sites, restore structural integrity and electrical conductivity, and prevent the expansion of local damage. Hydrogels with self‑healing features have the advantages of stable and reliable performance as well as longer service life.

Fig. 4.

The principle of Self-healing hydrogels. a Schematic diagram of self-healing hydrogels; b Schematic diagram illustrating the noncovalent bond; c Schematic diagram illustrating the covalent bond; d Schematic of self-healing process depending on hydrogen bond; e Macroscopic demonstration of the self-healing performance depending on hydrogen bond; f Schematic of self-healing process depending on dynamic borate ester bonds; g Macroscopic demonstration of the self-healing performance depending on dynamic borate ester bonds; h SEM image of self-healed hydrogel electrode/EL interlayer/Ni electrode. Black dashed boxes: cut/healed area. Heated at 50 °C for 24 h. Scale bar: 20 μm; i Schematic illustration of external stimuli-promoted hydrogel self-healing; j Macroscopic demonstration of external stimuli-promoted hydrogel self-healing

The advantages of non‑covalently cross‑linked hydrogels include short healing time and high healing efficiency. Among them, hydrogen bonds form reversible cross‑links between different components in hydrogels through the direct bonding of hydrogen atoms and highly electronegative atoms. Researchers fabricated a self-healing hydrogel with multiple dynamic hydrogen bonds between the hydroxyl groups of PVA and the carboxyl groups of AAc (Fig. 4d). The obtained hydrogel recovers 72% of its initial tensile strength within 12 h [110] (Fig. 4e). Although the rapid formation of hydrogen bonds improves the self‑healing speed of hydrogels, hydrogen bonds possess lower strength compared with ionic bonds and covalent interactions. As another self‑healing mechanism, ionic bonds are reversible electrostatic interactions between oppositely charged ions, which can be flexibly generated between ions or charged polymers. For example, researcher developed an oxidized‑cnf/PANI/PAA hydrogel, which exhibits self‑healing capability because Fe³⁺ forms reversible metal‑ligand coordination bonds with carboxyl groups in oxidized‑cnf and PAA chains [111]. Hydrophobic interaction refers to the spontaneous aggregation behavior of non‑polar hydrophobic groups in aqueous solutions and can be controlled by adjusting the shape and number of hydrophobic groups. When the hydrogel is damaged or cut, polymer chains within hydrophobic regions undergo self‑reorganisation around the affected area. This reorganisation generates attractive forces between adjacent hydrophobic regions and facilitates the self‑healing process of damaged areas. As another autonomous self‑healing strategy, hydrogels can achieve self‑healing via cross‑linking by reversible chemical bonds. These bonds deliver higher bond strength compared with non‑covalent bonds. At present, dynamic covalent bonds (including Schiff base bonds, disulfide bonds and boron/boronate ester bonds) and metal chelation/coordination bonds have been effectively utilized to fabricate self‑healing hydrogels. These covalent bonds are formed via reversible coupling, enabling hydrogels to maintain a dynamic equilibrium between bond cleavage and reformation. Researchers fabricated a hydrogel with the formulation of PVA-B/SA-C/LiCl/Gly-PVA. The dynamic borate ester bonds formed with borax can undergo spontaneous dissociation and reconstruction at room temperature [112] (Fig. 4f-h).

Besides autonomous self‑healing, the self‑healing of CHs can also be triggered by external stimuli (such as light and heat). For example, Xu et al. developed a rapidly self‑healing CH using the classic DN structure (Fig. 4i), which involves Sb₂S₃ nanorods coated with da‑modified PPy in a PVA/poly(N‑(2‑hydroxyethyl)acrylamide) matrix in a water‑glycerol solution (Fig. 4j). Notably, the modified nanorods exhibit photothermal conversion properties. Upon exposure to near‑infrared (NIR) light, the targeted damaged areas within the hydrogel can achieve rapid, in‑situ self‑healing.

Electrical properties

Conductivity, as the core function of conductive hydrogels, is the foundation for their translational application. Conductive hydrogels are usually composed of conductive materials and polymer networks, where the polymer network acts as a structural scaffold and the conductive materials endow the gel with conductivity [113, 114]. Methods to enhance the conductivity include doping with highly conductive fillers and optimizing conductive pathways [95, 115]. Conductive polymers can be doped to generate more charge carriers on the polymer chains, thereby further enhancing their conductivity. However, excessive doping reduces the dispersibility of fillers and impairs the mechanical properties of hydrogels. Optimizing conductive pathways improves conductivity by inducing the directional alignment of conductive components to construct efficient transport networks, thereby avoiding the issues of long transport paths in randomly dispersed networks. Researchers fabricated conductive polymer hydrogels by doping polypyrrole (PPy) and crosslinking it with TA and Fe³⁺ ions during in-situ polymerization [100]. In CPHs, PPy not only serves as the structural scaffold but also acts as the conductive component. Fe³⁺ facilitates the polymerization of pyrrole via oxidation and further coordinates with TA to act as an ionic crosslinker to form an ionic crosslinked network. Thus, this conductive gel demonstrates enhanced electronic conductivity (0.05–0.18 S·cm^-1). In addition, the conductivity and mechanical properties of conductive polymer hydrogels can be modulated by precisely controlling the concentration of dopants and crosslinkers (Fig. 5a-b). Researchers prepared hydrogels with effectively improved conductivity by forming conductive pathways through directional alignment. Mechanical stretching was used to induce the alignment of silver nanowires (AgNWs), deformable liquid metal (LM), and PVA, enhancing the connectivity between conductive fillers and further improving the conductivity of the hydrogel (0.24 S·cm^-1) [116]. Meanwhile, the effective combination of PVA, AgNWs, and LM effectively alleviates stress concentration at crack tips, significantly improving the mechanical properties of the hydrogels.

Fig. 5.

Strategies for electrical property improvement. a Schematics of the synthesis of percolated polypyrrole (PPy) network guided by the nanofiber template; b Conductivity demonstration of conductive nanofiber hydrogels (CNHs) via LED array (1 cm); serpentine pattern on ANF-PVA before/after 80% elongation. Scale bars: 1 cm; c Schematics of PEDOT: AlgS foams; d Conductivity enhancement of alginate with PEDOT fillers (7.1 × 10 –4 S·cm-1, n = 3); e Sample of PEDOT: AlgS; f PEDOT dopant effect on alginate ink impedance (inset: Nyquist plots); g Capacitive charge injection mechanism; h Faradaic charge injection mechanism

In addition to the directional alignment of conductive pathways, improving the dispersibility of conductive fillers is also an effective strategy for enhancing electrical conductivity. By strengthening the affinity between conductive fillers and solvents, agglomeration can be effectively suppressed to form a compact conductive network. Researchers replaced PSS with natural hydrophilic dopants, which boosted the dispersion concentration by 5-fold, enabled the loading of conductive fillers at a higher content, and synergistically reduced charge transport resistance. Compared with conventional PEDOT: PSS, the electrical conductivity of the hydrogel was improved by approximately 20-fold [117] (Fig. 5c-e).

Beyond bulk electrical conductivity, a suite of interfacial electrochemical properties dictates the practical performance of conductive hydrogels at biological interfaces, including frequency-dependent interfacial impedance, charge storage capacity, charge injection capacity, and electrochemical stability. Interfacial impedance is a key parameter that determines the signal‑to‑noise ratio and fidelity of signal acquisition as well as the occurrence of motion artifacts [118]. Its frequency dependence originates from two parallel acting mechanisms: capacitive transduction and Faradaic transduction. As the signal frequency varies, the contribution ratio of the two mechanisms changes, causing the magnitude and phase of interfacial impedance to vary regularly with frequency. Generally, various conductive hydrogels exhibit decreasing impedance as frequency increases.

ICHs are purely ionic conductors. Ion‑electron transduction can only be achieved via the electric double layer at the interface in contact with metal electrodes. Accordingly, they exhibit high low‑frequency interfacial impedance and prominent frequency dependence, which degrades the signal‑to‑noise ratio for acquiring low‑frequency physiological signals such as electrocardiograms and electroencephalograms [119]. In contrast, electronically conductive hydrogels feature conjugated polymers as their core component. They enable electron transport and support Faradaic processes through reversible redox doping/dedoping reactions on conjugated chains. They possess lower overall interfacial impedance and can maintain low impedance over a broad frequency range, making them suitable for both low‑frequency biological signal acquisition and high‑frequency electrical stimulation [120].

Charge‑injection capacity refers to the maximum amount of charge that can be injected per unit electrode area without triggering irreversible electrochemical reactions or tissue damage, and it determines the implementation efficiency of bioelectrical stimulation [121]. The charge injection of ion‑conductive hydrogels relies on electric‑double‑layer charging‑discharging and directional ion migration, which is limited by ion concentration, mobility and effective interfacial area, leading to low charge‑injection capacity and poor electrical stimulation efficiency. Electron‑conductive hydrogels based on conductive polymers accomplish charge injection through the Faradaic process of reversible doping‑dedoping, and can deliver a large amount of charge at relatively low voltages [122, 123].

Electrochemical stability under long‑term working conditions is critical for wearable and implantable bioelectronic devices [124]. Ionic‑conductive hydrogels achieve electrical conduction via ion migration without undergoing redox chemical reactions (Fig. 5g). The materials barely degrade during long‑term cycling and exhibit excellent electrochemical stability, making them more suitable for passive monitoring devices requiring long‑term implantation. Electron‑conductive hydrogels based on conducting polymers are prone to dopant ion loss and microstructure collapse after repeated electrochemical cycling, accompanied by gradual declines in electrical conductivity and charge capacity (Fig. 5h). Their insufficient long‑term stability constitutes a major bottleneck for their application in long‑term implantable devices [125, 126].

Mitigating the foreign-body response and introducing stimulus responsiveness

In addition to the aforementioned conformal contact, stability, and electrical conductivity, there are something worthy of attention, including enhanced biocompatibility and stimulus responsiveness. Improving biocompatibility via chemical modification of materials suppresses the recognition and activation of macrophages towards implants, alleviates local inflammatory responses, and inhibits collagen deposition and fibrotic encapsulation. Reducing the foreign body response suppresses the recognition and activation of macrophages toward implants, alleviates local inflammation, and inhibits collagen deposition and fibrotic encapsulation via material chemical modification [127]. Typical strategies to mitigate the foreign body reaction include chemical group modification, the introduction of anti-inflammatory factors, and biomimetic coating of hydrogels [128]. Researchers fabricated a transparent conductive hydrogel with a double-crosslinked network, featuring both reparability and biocompatibility [129]. In vitro assays showed that the survival rate of HEK293 cells exceeded 84.39%, and no inflammatory response was observed 14 days after in vivo implantation. Researchers fabricated a real-time glucose-interactive dynamic hydrogel dressing with favorable biocompatibility. The hydrogel system was co-cultured with human umbilical vein endothelial cells (HUVECs) for 5 days; Live/Dead staining showed no increase in dead cells, and the CCK-8 assay demonstrated that the cell viability was consistent with that of the blank control. Researchers also developed a microenvironment-feedback hydrogel as a living dressing biomaterial. Live/dead cell staining assays demonstrated that when HUVECs were co-cultured with the hydrogel for 1 and 3 days, the proportion of live cells was close to that of the control group (without hydrogel) and the number of dead cells was negligible, indicating no obvious cytotoxicity of the hydrogel. CCK-8 assays revealed that after 1, 3, and 5 days of culture, there was no significant difference in cell viability between the hydrogel group and the control group (P > 0.05), further confirming its favorable cytocompatibility [130] (Fig. 6a-c).

Fig. 6.

Strategies for mitigating foreign body reaction and enhancing stimulus responsiveness. a Live/dead staining of HUVECs ± hydrogel (Day 1/3). Scale bar: 50 μm; b HUVEC viability ± OSA-GEL@GC (Day 1/3/5, triplicate, n = 3); hemolysis test (ddH₂O/PBS/hydrogel, triplicate, n = 3); c Images of transwell experiment of HUVECs treated with or without hydrogel group. Scale bar: 50 μm; d Photograph of the two eutectogels; e Relative resistance changes during 10,000 consecutive finger pressing with the pressing degree is 1700 − 1650. The insets are the first 10 cycles and the final 10 cycles

Stimulus responsiveness refers to the ability of a hydrogel, as a three-dimensional hydrophilic network, to respond to external signals (such as temperature, pH, chemicals, light, electric field, and shear stress) and thus change its swelling behavior or physical properties [131–134]. In terms of applications, sensitivity determines whether hydrogels can detect changes and convert them into signals. Various strategies have been developed to improve the stimulus responsiveness of conductive hydrogels [135]. These mainly involve the design of dynamic reversible networks, the introduction of stimuli-responsive components, and the hybridization of functional nanomaterials. Researchers synthesized peptide-enhanced eutectic gels by introducing the α-helical “molecular spring” structure into deep eutectic solvents (Fig. 6d). The mechanical strength of the gels is improved significantly while maintains highly resilient simultaneously. Therefore, the sensor possesses excellent stability. One can see both the peak values and the curve shapes maintain the same even after 10,000 bending cycles (Fig. 6e). Benefiting from the rational engineering strategies toward conformal contact, stability, conductivity, biocompatibility and stimulus responsiveness, conductive hydrogels have exhibited extensive adaptability and unlocked a wealth of advanced applications in flexible bioelectronics.

In summary, Table 1 provides an overview of representative conductive hydrogels recently developed for bioelectronic interface applications. The summarized parameters include mechanical properties (elastic modulus, tensile strain and durability), electrical properties (conductive mechanism, electrical conductivity), and interfacial/bio-related properties (adhesion energy and biocompatibility).

Table 1.

Summary of representative conductive hydrogels for bioelectronic interfaces and their key performance parameters

Material	Type	Conductivity (S m− 1)	Modulus	Strain	Adhesion	Durability	Biological compatibility	Application	Ref
Na+	ICH	4.03	5.2 MPa	235%	-	2000 cycles	-	Nanogenerator Sensor	[30]
Al+	ICH	-	0.29 MPa	696%	-	-	-	Sensor	[31]
Li	ICH	1.71	≥ 19.2 kPa	1348%	-	-	-	Sensor	[32]
PEDOT: PSS	ECH	7400	3.9 MPa	40%	-	-	√	Stimulating electrode	[136]
PANI	ECH	4.26	1.37 MPa	110%	-	-	-	Sensor	[137]
PPy	ECH	8000	9.4 MPa	36%	-	-	√	Sensor	[100]
AgNPs	ECH	0.41	261 kPa	515%	-	-	-	Sensor	[44]
AgNWs	ECH	90,000	240 kPa	480%	-	2850 cycles	√	Electrode Flexible circuit	[49]
EGaIn	ECH	0.02	6.24 kPa	2000%	960 kPa	-	-	Sensor	[138]
CNTs	ECH	0.145	22.7 kPa	-	6.03 kPa	-	-	Sensor Nanogenerator	[52]
rGO	ECH	3.63	250 kPa	350%	-	50 cycles	√	Sensor Wound dressing	[59]
	ECH	8	75 kPa	3400%	30 kPa	-	√	Sensor Implanted electronic devices	[57]
MXene	ECH	0.33	42.5 kPa	729%	-	-	-	Flexible electronics	[139]
PEDOT: PSS; AlCl	Hybrid CH	3800	27 kPa	45%	-	-	-	Flexible electronics	[140]

Emerging applications of conductive hydrogels in bioelectronic interfaces

Wearable flexible electronic devices

The application of multifunctional conductive hydrogels is currently mainly focused on the design and development of wearable flexible electronic devices [141–144]. Conductive hydrogels can synergistically integrate multiple properties, including conductivity, mechanical compliance, self-healing capability, and adhesion. These features endow wearable sensors with numerous application advantages, such as high detection sensitivity and measurement accuracy, stable operational reliability, long-term service stability, and stimulus-responsive intelligence [11, 145–147]. For instance, to address the stability issues of traditional hydrogel electrodes caused by mechanical brittleness, easy dehydration, and freezing, researchers developed a robust hydrogel electrode. This hydrogel electrode achieves excellent antidehydration and antifreezing properties(-25℃) through the component synergy of a glycerol-water binary solvent and lithium chloride (LiCl), combined with a gelatin cross-linked network [148]. It enables long-term continuous monitoring and provides an effective strategy for improving the stability of wearable flexible electronic devices(Fig. 7a). In recent years, multimodal integrated sensors have become a research hotspot in this field [145, 149–151]. The evolution from single-parameter detection to multimodal collaborative sensing represents a fundamental paradigm shift in sensing technology [152, 153]. Researchers integrate dual sensing capabilities of strain and temperature into a conductive hydrogel, which can be conformally attached to the human body to accurately monitor motor behaviors and distinguish physiological signals [154] (Fig. 7b-c). This application enables the detection of diverse external stimuli and physiological signals, improving the comprehensiveness and accuracy of health monitoring. This enables wearable flexible electronics to move beyond mere monitoring toward integrated therapy and recognition.

Fig. 7.

Applications of wearable hydrogels. a 8-day ambulatory ECG monitoring via ultrathin hydrogel electrodes (inset: ECG waveforms during daily activities); b temperature sensor; c temperature sensor; d Schematic illustration of the electronic bandage based on high-strength fracture-resistant ionogels (HSFRIs); e Full-thickness wound healing (Day 0/1/3/6/9). Scale bar: 2 mm

In addition, conductive hydrogels hold promise in biomedical and wearable healthcare applications [155]. Compared with traditional clinical wound dressings and conventional bandages, the electronic bandages fabricated by researchers exhibit advantages. For instance, ionic conductivity, drug-loading capacity, and the conformal contact of conductive hydrogels render them an ideal matrix for electronic bandages. This conductivity enables a synergistic effect with electrical stimulation to accelerate cell proliferation, collagen deposition, and vascular regeneration [156]. The optimal e-bandage (high-strength and fracture-resistant ionogels-erythromycin-electrical stimulation, HSFRIs-EM-ES) achieved a wound contraction rate of 96.84% on day 9, which was 21.1% points higher than the blank control group (75.74%) [157] (Fig. 7d-e). This endows the electronic bandage with unique electrotherapeutic and healing-promoting functions unavailable in traditional bandages.

Implantable hydrogels

Compared with research on wearable devices, implantable flexible electronics are less studied at present [146, 158–160]. However, they possess broader prospects in the long term. This is because conductive hydrogels usually possess high water content and low modulus, which make them physically similar to the extracellular matrix (ECM) [161–164]. Consequently, they can reduce immune rejection and inflammatory responses, laying a solid foundation for long-term implantation [165–167]. For implantable flexible brain-computer interfaces, researchers have developed a hydrogel neural interface. Compared with metallic electrodes, this device triggers fewer immunochemical reactions and enables long-term in vivo application. Electrical stimulation enabled by conductive hydrogels can also be applied to the field of neural regeneration [168]. For example, researchers fabricated a conductive hydrogel capable of delivering electrical stimulation and recording electrical signals (Fig. 8a-b). This hydrogel can help restore electrical signal conduction between damaged nerves and muscles [169]. Compared with existing rigid implantable electrodes in practice, this hydrogel can suppress foreign body inflammation and fibrotic encapsulation. This improves the long-term stability of the implant interface and therapeutic efficacy. However, the clinical therapeutic efficacy remains to be further validated.

Fig. 8.

Applications of implantable hydrogels. a Wireless deep brain stimulation (DBS) setup: injectable conductive hydrogel (ICH) in STN; neural activity recorded in M1/GPi; b Representative images show in situ gelation of ICHs in tissue-mimicking agarose gel over time. Scale bar: 5 mm; c Schematic illustration of hydrogel cardiac patches for myocardial infarction monitoring and repair; d Remote measurement of ECG signals in rats via DTMS and a Bluetooth module; e Recorded ECG signals

In addition to applications in brain-computer interfaces and spinal cord stimulation, conductive hydrogels can also be utilized for monitoring and treatment in the cardiovascular field [170, 171]. Researchers proposed a microchannel hydrogel suture for bidirectional communication (Fig. 8c-e). It acts as a cardiac patch to repair infarcted myocardium. It can also remotely monitor the heart and deliver drugs on demand [9]. This realizes the integration of diagnosis and treatment in implantable flexible devices. Compared with common electrocardiogram (ECG) monitoring, this method has better performance. It provides convenience while enabling drug delivery, and hold promise for the integration of diagnosis and therapy in the future.

Key challenges and future perspectives and research directions

Despite satisfactory progress of conductive hydrogels in the field of bioelectronic interfaces, their translation from laboratory prototypes to clinical and industrial applications still faces multidimensional bottlenecks in material design, interfacial interactions, fabrication processes, biosafety, and other aspects. Based on the current research status, the core challenges and future research directions are summarized as follows:

Challenges

First and foremost, conductive hydrogels face a pervasive trade-off dilemma in the synergistic optimization of multiple key properties [172]. Most hydrogels still fail to break through such performance bottlenecks. Specifically, high mechanical strength or toughness usually reduces water content and weakens electrical conduction. High conductivity often requires high loading of conductive fillers, leading to increased modulus and poor tissue compliance. Functional modifications (e.g., adhesion and self-healing) tend to sacrifice structural stability or biocompatibility. Thus, the simultaneous and balanced optimization of softness, conductivity, adhesiveness, stability, and biocompatibility remains the core challenge in this field.

Meanwhile, limitations exist in fabrication processes and integrated manufacturing. High-performance conductive hydrogels are mainly synthesized through methods (solvent exchange, in-situ polymerization, freeze-casting, etc.). However, in situ polymerization may pose limitations on scalability and cost-effectiveness in industrial applications because of the complexity of reaction conditions and extended preparation times. Achieving uniform dispersion and solubility of nanomaterials in hydrogels remains a significant challenge. The process of freeze-casting can be time-consuming and requires precise control over freezing and thawing conditions to ensure consistency and reproducibility. To ensure uniform and defect-free layers, Layer-by-layer assembly requires meticulous control of deposition conditions, which may add complexity to the fabrication process. The precursors in sol–gel process may introduce trace impurities that can significantly influence the properties of the synthesized composite hydrogel, necessitating stringent quality control measures. These limitations make them unsuitable for large-scale and continuous industrial production. Second, standardized evaluation systems and biosafety certifications are still lacking. No unified standards have been established for characterizing mechanical, conductive and interfacial stability properties [173]. This makes it hard to horizontally compare performance among different studies. For implantable applications, long-term metabolic pathways, degradation toxicity and chronic inflammatory responses lack sufficient in vivo validation, especially for nanomaterial-containing composite hydrogels. The insufficient biosafety verification slows down the certification process of medical devices. Lastly, the adaptability of conductive hydrogels to extreme environments needs to be enhanced. Their inherent drawbacks (easy dehydration, drying and freezing) potentially limit their further practical applications [174].

Outlook

Overcoming the trade-off dilemma requires careful design strategies that balance these properties while ensuring long-term reliability. Multicomponent and hierarchically synergistic networks can be developed via molecular engineering and nanotechnology to balance diverse performance requirements. This strategy achieves high conductivity, high toughness and tissue-matched softness simultaneously. Combined with 3D bioprinting and microfluidics, customized fabrication of hydrogel devices can be realized [172]. It is urgent to establish standardized evaluation frameworks and unified testing criteria for related systems as early as possible. Long-term large-animal in vivo experiments are also indispensable to systematically validate the biosafety and degradation behavior of functional hydrogels. Such fundamental endeavors will greatly advance the translational progress of hydrogel-based bioelectronics, facilitating their transformation from laboratory research to practical clinical trials.

Moreover, the function integration including diagnosis and treatment represents a promising direction. In parallel, self-powered wearable electronic systems can be constructed by leveraging the adequate properties of these hydrogel materials. With the assistance of artificial intelligence algorithms, it is feasible to realize intelligent physiological signal recognition, precise disease diagnosis, and on-demand therapy including electrical stimulation and controlled drug delivery. Ultimately, a fully functional intelligent bioelectronic platform can be well constructed to support intelligent healthcare applications [173].

Conclusions

With tissue-like mechanical compliance, tunable electrical conductivity, and favorable biocompatibility, conductive hydrogels become core bridging materials for next-generation bioelectronic interfaces. They effectively reconcile the mechanical and electrical property mismatch between biological tissues and artificial electronic hardware, laying a solid foundation for the advancement of flexible bioelectronics. This review systematically summarizes the fundamental design strategies of conductive hydrogels for bioelectronic interface applications. It focuses on the mechanism of conductive hydrogel and elucidates the core properties, advantages, and existing research results.

Focusing on the demand for efficient adaptation at the hydrogel-tissue interface, this paper elaborates on the core engineering strategies to improve the practical performance of conductive hydrogels. These strategies include three key approaches to achieve conformal contact: modulation of material mechanical compliance, ultrathin design, and adhesion optimization. Toughness enhancement and self-healing functionalization are employed to strengthen structural stability and service reliability. Conductivity optimization is utilized to guarantee efficient signal transmission. Material modification strategies are proposed to reduce foreign body reactions and improve biocompatibility. Functionalization design is also adopted to endow materials with stimulus-responsive properties. Together, these approaches establish a comprehensive performance optimization system. At the application level, this review highlights the breakthrough progress of conductive hydrogels in wearable and implantable flexible electronic devices. In the wearable field, conductive hydrogels have been applied in electronic bandages, multimodal physiological monitoring sensors, sign language translation human-machine interaction systems, and other scenarios. These researches realize the health monitoring, wound therapy, and intelligent interaction. In the field of implantable medicine, existing studies report the application of conductive hydrogels in brain–computer interfaces, spinal cord stimulation, and cardiovascular monitoring in animal models. These systems hold promise for translation into clinical therapeutic approaches in the future.

This review clarifies the key to developing high-performance conductive hydrogels for bioelectronic interfaces. The core lies in achieving synergistic optimization of multiple functions and breaking the trade-off bottlenecks among conformal contact, stability, electrical conductivity, and biocompatibility. Despite the substantial progress achieved, this field still confronts prominent challenges including the difficulty in coordinated regulation of multiple performance parameters, underdeveloped large-scale fabrication techniques, the absence of standardized evaluation to chronic biosafety, and sterilization effect. Moreover, the long-term degradation, electrochemical stability and manufacturing reproducibility of conductive hydrogel devices require convincing results. Strict controlled clinical studies have not been conducted in target patient populations for specific disease scenarios such as diabetes and cardiovascular diseases, nor have core clinical indicators including the diagnostic accuracy, specificity, and positive/negative predictive values of the sensor been verified using gold‑standard methods. Nevertheless, continuous technological innovation will further accelerate the evolution of this research direction. Driven by advances in material molecular engineering, interface engineering and intelligent system integration, conductive hydrogels will continue to fuel the development of wearable and implantable electronic devices. Serving as an indispensable link bridging biological systems and artificial electronics, conductive hydrogels retain profound research significance and promising application prospects for future clinical and intelligent healthcare innovation.

Author contributions

H.C., T.L., and X.C. contributed equally to this work. M.S., W.Z., H.Z., and Y. J. conceived the work. H.C., T.L., X.C., Z.C., and H.T. organized the content and prepared the draft manuscript. M.S. and W.Z. supervised the work. All authors contributed to the writing and revision of the manuscript.

Funding

The authors are grateful for financial support from Brain-Gain Plan of New Chongqing (Young Elite Program) and the State Key Laboratory of Precision Welding & Joining of Materials and Structures (MSWJ-24M11).

Data availability

No datasets were generated or analysed during the current study.

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All of the authors declare that this is a review article and did not require research involving human participants; thus, no informed consent was needed.

Competing interests

The authors declare no competing interests.