At organ interfaces that require high-density electrode arrays and long-term stable recording and stimulation, the structural layer must both support precise micro- and nanopatterns and maintain its geometry and mechanical properties over years^[26]. To meet these demands, a class of high-modulus, non-degradable polymer films has become the standard structural material set, including Parylene-C, polyimide (PI), polyethylene naphthalate (PEN), and liquid crystal polymers (LCPs) [Figure 3A]^[27-31]. These materials typically exhibit Young’s moduli in the gigapascal range, far above those of soft tissues, but when the film thickness is reduced from tens of micrometres to a few micrometres or even sub-micrometre levels, their bending stiffness decreases substantially^[32]. This thickness scaling allows them to conform to curved organ surfaces such as the cerebral cortex, retina, or peripheral nerves without pronounced wrinkling or structural instability, while still preserving in-plane dimensional stability and pattern fidelity.

These polymers are highly compatible with established microelectronic fabrication processes and can accommodate standard steps such as photolithography, metal interconnect patterning, and multilayer via formation. As a result, they are widely used as structural substrates for microelectrocorticography (μECoG) arrays, retinal prostheses, and cuff electrodes targeting the auditory and peripheral nerves. Ultrathin parylene-C films can further serve as transferable and biocompatible carrier substrates because of their excellent conformability, low-temperature processability, and water resistance^[33,34]. For example, Cho et al. developed a textile-integrated OLED platform in which a thin parylene-C film was first formed on a guide glass and then transferred onto textile after thermal annealing, serving as a self-supporting planarization layer that replicated the smooth glass surface while preserving textile flexibility. This strategy enabled stable device fabrication and operation on rough, deformable textile substrates [Figure 3B]^[35]. At the same time, these materials provide good electrical insulation and moisture barrier performance, and can therefore serve as the primary encapsulation layer of the device. For example, parylene-C deposited by chemical vapor deposition can form dense, conformal, and nearly pinhole-free coatings with stable dielectric properties, making it one of the most widely used insulating and encapsulation materials for implantable neural electrodes and cardiac pacing leads [Figure 3C]^[36]. LCPs combine low water vapor transmission with the ability to be thermoformed and laminated, making them suitable for constructing hybrid rigid–flex encapsulation shells in intracranial or cochlear implants where higher mechanical robustness is required^[37].

Overall, non-degradable high-modulus films represented by Parylene-C, PI, PEN and LCPs remain the predominant choice for organ interfaces that experience relatively small surface strain but demand high spatial resolution and long-term electrical stability, such as the cerebral cortex, retina and peripheral nerves^[38-40]. Looking ahead, future efforts are likely to focus on retaining the mature processability and encapsulation reliability of these materials while introducing softer interlayers or biohybrid architectures to improve tissue integration. For example, Shi et al. developed a biohybrid multilayer interface combining PI, Au, SU-8, and living hydrogel layers to improve tissue-level compliance and interfacial integration with soft biological tissues^[41].

For highly deformable organs and tissues such as the skin, heart, and diaphragm, which experience substantial, multidirectional deformation during routine motion, an overly rigid structural layer can readily lead to electrode delamination, interconnect failure, or localized tissue compression. To address this, low-modulus elastomeric substrates based on polydimethylsiloxane (PDMS), Ecoflex, polyurethane elastomers (PU) and thermoplastic polyurethane (TPU) have become widely used at high-deformation organ interfaces [Figure 3D]^[42]. These materials combine flexible polymer backbones with relatively low crosslinking density, allowing their Young’s modulus to be tuned into the 10³-10⁷ Pa range, while sustaining 50%-200% tensile strain without plastic damage. Shin et al. reported a stretchable multichannel sensor array, illustrating how mechanically compliant elastomeric substrates can be incorporated into bioelectronic systems to enable stable impedance and pH mapping under both static and dynamic conditions [Figure 3E]^[43]. This tissue-like mechanical compliance allows the device to conform and move with skin, cardiac, and diaphragmatic tissues during vigorous motion, respiration, and heartbeat, thereby reducing stiffness mismatch and improving interfacial stability.

With these substrates, geometric layout further determines the effective operating range and failure mode of stretchable platforms under high strain. A widely adopted strategy is to confine rigid or semi-rigid functional units to local low-strain zones and connect them to the surrounding elastomer via deformable interconnects. Representative strategies also exploit interfacial engineering between conductive layers and elastomeric or fibrous substrates. For example, conductive coatings can be stabilized through molecular locking, pre-strain-induced wrinkling, and confined interpenetrating nanofiber networks, thereby maintaining electrical continuity and mechanical compliance under repeated deformation^[44]. Such strain-engineered architectures enable graded dissipation of mechanical deformation around rigid functional components, thereby maintaining electrical continuity and signal stability during repeated stretching and release. At the material level, such interconnects are typically realized using metal thin films, metal meshes or embedded liquid-metal channels, all encapsulated within PDMS, Ecoflex or PU matrices^[45-47]. By tailoring interconnect geometry, including interconnect length, linewidth, and radius of curvature, externally applied strain can be redistributed into bending, twisting, and local rotation of the interconnects, thereby reducing strain concentration in functional electrode and chip regions. This strain-isolation strategy helps preserve interfacial contact, contact impedance, and signal amplitude during prolonged body motion or repeated organ contraction^[48,49].

From the standpoint of encapsulation and interfacial stability, elastomeric platforms are usually realized as a multilayer system, in which an elastic substrate is combined with an elastic encapsulation layer and tailored surface modification. Kim et al. developed a photo-patternable Ecoflex encapsulation strategy that enables process-compatible formation of patterned elastic encapsulation layers with multi-windows, thereby improving strain dissipation, electrical stability, and selective multi-analyte sensing in intrinsically stretchable wearable bioelectronics [Figure 3F]^[50]. The internal conductors and electrodes are first fully embedded in PDMS, PU or TPU to provide electrical insulation and to reduce stress concentrations at geometric discontinuities^[20,51,52]. For devices that are worn on the skin or exposed to sweat and body fluids, an additional thin fluoropolymer overlayer, such as polytetrafluoroethylene (PTFE), can be applied to the outer surface to reduce water and ion permeation while maintaining sufficient water vapor transmission^[53-55]. For devices adhering to the epicardium or diaphragm, the thickness and modulus of the encapsulation layer must also be balanced against the frictional interaction with pericardium or pleura, so as to avoid excessive shear during cardiac or respiratory cycles.

In many short-term or single-use implant applications, the ideal interface material should gradually degrade and be resorbed in vivo once monitoring or therapy is complete, thereby avoiding secondary removal surgery and reducing long-term foreign body reactions. With this goal in mind, protein-based materials such as silk fibroin and gelatin, together with synthetic polyesters such as polylactic acid (PLA) and poly(lactic-co-glycolic acid) (PLGA), have become representative systems for biodegradable structural layers [Figure 3G]^[56]. Protein materials are degraded enzymatically through cleavage of amide bonds, whereas polyesters undergo hydrolysis of ester bonds to generate lactic acid, glycolic acid and other small molecules that subsequently enter normal metabolic pathways and are cleared^[57]. By pre-embedding such cleavable chemical bonds in the polymer backbone, these materials can be programmed to gradually lose mechanical integrity and disappear within a defined time window under the combined action of water and enzymes, providing the chemical foundation for fully resorbable support and encapsulation layers^[58].

Accordingly, the central design challenge for biodegradable structural layers is to coordinate mechanical support lifetime, electrical stability and degradation behaviour within an appropriate range^[59]. For PLA and PLGA, effective support time can be extended from several weeks to several months by tuning crystallinity, molecular weight and the copolymer ratio of lactic to glycolic acid^[60]. Higher crystallinity and molecular weight slow water ingress and backbone scission, making them suitable for medium-term cardiac or neural stimulators, whereas a higher glycolic acid content usually accelerates hydrolysis and is better suited to postoperative short-term monitoring patches or temporary electrodes^[61,62]. Silk fibroin and gelatin can be adjusted in an analogous way by varying β-sheet content, crosslinking density or integration with inorganic fillers, allowing stiffness and degradation rate to be tuned while maintaining good cell compatibility^[63]. Beyond conventional films and bulk scaffolds, biodegradable protein matrices can also be engineered into hollow microfiber assemblies and organ-mimetic microsystems. For example, Tian et al. developed a microfiber-assembled endocrine pancreas-on-a-chip based on microfluidically spun GelMA/CaA hollow fibers, in which the biodegradable hollow microfibers were used to mimic vascular lumens and support material transport, and were integrated with a 3D pancreatic islet culture layer for islet culture and functional evaluation [Figure 3H]^[64]. Silk fibroin-based polymers can also be engineered into mechanically robust and ionically conductive fibers by incorporating ionic liquid components. Lu et al. reported a silk fibroin-based ionic hydrogel fiber composed of silk fibroin, ionic liquid, and glycerol, which exhibited high strength, large extensibility, and stable ionic conductivity, thereby enabling deformable conductive fibers for wearable bioelectronic and human–machine interface applications [Figure 3I]^[56]. In addition to structural substrates, biodegradable matrices can also be integrated with transient sensing interfaces and soft conductive layers. For example, agar- and hydrogel-based ionic interfaces combined with thin metal electrodes can provide temporary skin-compatible sensing platforms with minimized long-term foreign-body burden after use^[65]. Another approach integrates ultrathin bioresorbable inorganic oxides as transient insulating or encapsulation layers and hydrolyzable conductive polymers as temporary charge-transport components, enabling the required electrical functions to be maintained over a prescribed time frame before gradual resorption^[66,67].

At the level of organ interfaces, biodegradable and bioresorbable structural layers are particularly attractive in three types of scenarios. The first includes the heart and peripheral nerves in the postoperative or acute phase, where temporary electrophysiological interventions such as transient pacing, vagus nerve modulation or post-surgical rhythm monitoring are required. The second concerns wound beds, graft sites or remodelling soft tissues, where staged mechanical and electrophysiological monitoring or local electrical stimulation is needed during healing, but complete device disappearance is desired once tissue repair is complete^[68]. The third involves patient groups for whom device retrieval is difficult or undesirable, such as neonates or very elderly individuals with multiple comorbidities. Recent work has begun to introduce dynamic crosslinks, inorganic fillers and multilayer gradient architectures into biodegradable frameworks, more tightly coupling mechanical support, electrical stability and controlled resorption^[69]. These strategies aim to align the functional lifetime of the structural layer more precisely with organ healing timelines, while reducing issues such as swelling during degradation, fragment migration and local irritation, thereby opening broader materials and structural design space for single-use implants and resorbable physiological monitoring systems.

In real physiological environments, once a device surface is exposed to plasma, cerebrospinal fluid or intestinal fluid, plasma proteins, inflammatory cells and microbes begin to adsorb and accumulate within a very short time^[70]. This rapidly triggers inflammatory reactions, thrombosis and fibrotic encapsulation, which ultimately manifest as increased electrode impedance, reduced signal amplitude and, in severe cases, complete device failure^[71,72]. Consequently, beyond the structural and encapsulation layers, a thin outermost interfacial coating that directly contacts body fluids and is engineered to resist protein adsorption, cell adhesion and even bacterial colonization has become a critical element for achieving long-term stable organ–device interfaces^[73].

One of the earliest systematically explored strategies is the construction of a hydrated, hydrophilic layer. A representative approach is to graft or coat polyethylene glycol (PEG) or related hydrophilic polymers onto the device surface so that, in aqueous environments, a dense, nanometre-scale hydration layer is formed^[74,75]. This hydrated shell introduces steric hindrance and a high energetic penalty for dehydration, thereby suppressing the residence of proteins and cells at the interface^[76]. Because PEG is susceptible to oxidation and hydrolysis under long-term implantation conditions, recent work has increasingly shifted toward zwitterionic polymer systems, such as coatings bearing sulfobetaine or carboxybetaine side chains^[77]. These materials carry both positive and negative charges on the same side chain, enabling the formation of more stable and tightly bound hydration shells. As a result, they exhibit more durable resistance to protein adsorption and cell adhesion in complex media such as plasma and cerebrospinal fluid, and have been applied to the surface modification of neural electrodes, vascular stents and catheters, where animal studies have shown reduced inflammation and attenuated glial scarring.

A second line of development draws inspiration from naturally “slippery” interfaces such as those of pitcher plants, through the creation of slippery liquid-infused porous surfaces (SLIPS)^[78]. In these systems, a porous fluoropolymer or micro/nano-rough substrate serves as a scaffold in which a thin layer of medical lubricant is retained within pores or topographical features, forming a continuous, stable and self-healing liquid interface that offers virtually no effective anchoring sites for bacteria, proteins or cells. Multiple animal models have demonstrated that SLIPS coatings can markedly suppress bacterial biofilm formation and fibrotic encapsulation, making them attractive for high-contamination-risk devices such as urinary catheters, vascular implants and GI systems. Unlike hydrophilic coatings that rely primarily on chemical repulsion, SLIPS function mainly through physical dewetting and re-spreading of the infused lubricant, allowing damaged or abraded regions to recover their antifouling properties from the lubricant reservoir.

For brain, retinal and peripheral nerve electrodes, interfacial coatings must not only mitigate inflammation and glial scar formation but also preserve low impedance and ionic transport^[79]. In intravascular or intestinal settings, additional requirements include antithrombotic and antibacterial performance and compatibility with fast-flowing fluids. Rationally matching these interfacial coatings with the non-degradable, elastic or biodegradable structural–encapsulation systems discussed above offers a route to constructing integrated interfaces that combine mechanical matching, chemical stability and biological tolerance across diverse organ environments, thereby providing a robust foundation for the long-term reliable operation of conductive layers and active electronic components^[80].

OMIECs are soft polymeric materials that transport both electronic and ionic charges while remaining mechanically compatible with biological tissues^[91,92]. They are typically built from polymers with π-conjugated backbones combined with hydrophilic side chains or polyelectrolyte components, giving moduli in the MPa range, good bendability, and compatibility with solution processing^[93,94]. Unlike purely electronic organic semiconductors, OMIECs undergo bulk doping in aqueous electrolytes so that charge is stored throughout the material volume^[83]. This volumetric electrochemical coupling is particularly relevant at soft organ interfaces, where tissue impedance is dynamic and signal amplitudes are often small, requiring stable, low-impedance charge exchange rather than purely interfacial capacitive coupling^[43].

Their mixed conduction can be viewed as the cooperation of electronic and ionic pathways within a single network^[85]. Electrons or holes move along the π-conjugated backbone through locally ordered domains, with conductivity governed by backbone structure and doping level^[95,96]. For representative poly(3,4-ethylenedioxythiophene):poly(styrenesulfonate) (PEDOT:PSS)-based OMIECs, electronic conductivity can range from below 1 S·cm^-1 in pristine or structurally disordered films to 10²-10³ S·cm^-1 after secondary doping, acid treatment, or structural optimization. Ions migrate within hydrophilic side chains or polyelectrolyte phases and compensate the charges created on the backbone during oxidation or reduction. Because this electron–ion compensation occurs throughout the bulk rather than at a confined interface, OMIECs exhibit high volumetric capacitance, typically on the order of tens to hundreds of F/cm³ for PEDOT-based systems, enabling high organic electrochemical transistor (OECT) transconductance at low operating voltages as well as low-impedance, high-charge-injection electrode interfaces. Such characteristics are advantageous in organ-specific bioelectronic chips that must operate under strict electrochemical safety limits while maintaining efficient coupling to excitable tissues such as brain and myocardium. Recent work has shown that the electrochemical doping rate in OMIECs strongly depends on film microstructure, film morphology, and electrolyte environment, which in turn determine device response speed and operational stability^[97,98]. These kinetic factors are particularly important in neural and cardiac electrophysiology, where millisecond-scale temporal resolution is required^[83,99].

Recent studies further underscore the importance of electrochemical kinetics in OMIEC-based organ interfaces. Keene et al. showed that electrochemical doping in conjugated polymers can be hole-limited, highlighting that OMIEC response speed depends not only on ion transport but also on electronic charge propagation within the polymer phase^[100]. For organ-specific bioelectronic chips, such behavior directly links material-level transport dynamics to system-level bandwidth and energy constraints. Beyond transport kinetics, recent material-design strategies have further expanded the utility of OMIECs in soft biointerfaces. Montazerian et al. showed that replacing conventional hydrophobic PSS dopants with hydrophilic biomacromolecular AlgS dopants in PEDOT:AlgS improves aqueous dispersibility, molecular degradability, and ionic integration with hydrogel matrices, thereby enabling injectable and 3D-printable OMIEC-based bioelectronics [Figure 4A]^[101]. Together, these findings shift the focus from purely material characterization toward mechanism-informed optimization of kinetics, stability, and energy efficiency in physiological environments.

Overall, OMIECs represent a commonly adopted conductive-layer strategy in organ-specific bioelectronic chips, enabling low-impedance electrochemical coupling at wet, soft, and dynamically active tissue interfaces under low operating voltages^[102,103]. Their mixed ionic–electronic transport characteristics make them particularly suitable for high-density neural interfaces in the brain and retina, as well as conformal sensing and stimulation devices deployed on mechanically deformable tissues such as myocardium and intestinal mucosa^[102].

Conductive hydrogels are soft, water-rich polymer networks that are engineered to carry electrical signals while remaining mechanically similar to biological tissues. They are typically formed by three-dimensional crosslinking of natural or synthetic polymers such as polyacrylamide, alginate, or gelatin, which creates a stable, highly hydrated network^[104]. By further introducing ionic or electronic conductive components into this matrix, conductive hydrogels can preserve softness and tissue affinity and at the same time support reliable electrical signal transmission^[105-107].

Their electrical behavior is supported by two complementary mechanisms. First, the abundant water phase inside the hydrogel provides pathways for mobile ions, giving rise to ionic conduction^[108,109]. Second, adding conductive elements such as metal nanowires, carbon nanotubes, graphene, or conducting polymers like PEDOT:PSS allows the formation of continuous electronic or mixed ion–electron pathways within the network^[110-113]. Depending on the conductive filler, polymer composition, and percolation structure, conductive hydrogels can span several orders of magnitude in conductivity, from predominantly ion-conducting networks with conductivities typically in the 10^-3-10^-1 S·cm^-1 range to electronically conductive composite or conductive-polymer hydrogels with conductivities reaching 1-10² S·cm^-1 or higher^[114]. Dynamic crosslinks based on hydrogen bonding, metal–ligand coordination, or catechol chemistry can break and reform during stretching, bending, or repeated loading. These reversible bonds help maintain the integrity of the conductive pathways and endow the material with self-healing and network reconfiguration^[115-117]. Recent work has shown that supramolecular poly(ionic) networks can simultaneously achieve ionic conductivities up to 0.1 S·cm^-1 and stretchabilities exceeding 1,500%, illustrating how dynamic crosslinking can reconcile efficient ionic transport with tissue-like deformability in multilayer conductive hydrogels [Figure 4B]^[24]. For example, metal-catechol coordination conductive hydrogels have been shown to recover both mechanical strength and electrical continuity after physical fracture, highlighting the role of dynamic bonding in sustaining cyclic stability^[118]. Such dynamic bonding is particularly beneficial for organ interfaces subjected to cyclic deformation, such as myocardium and diaphragm.

Recent advances continue to expand the functional scope of conductive hydrogels toward environments and use-cases previously inaccessible to soft materials. One major direction focuses on improving environmental robustness, as exemplified by Zhang et al., who introduced a “hydro-locking” strategy that stabilizes water within double-network hydrogels, maintaining softness and conductivity from -115 to 143 °C for extreme-condition sensing^[119]. Although originally demonstrated for extreme-condition sensing, such hydration-stabilization strategies may also improve long-term stability of implanted organ interfaces. A complementary direction integrates energy-harvesting and therapeutic functions, demonstrated by Xin et al., who developed a Fe²⁺/Fe³⁺-alginate thermogalvanic dressing capable of converting wound-site temperature gradients into therapeutic electrical stimulation while enabling real-time monitoring^[120]. This multifunctional integration hints at future organ-resident patches capable of simultaneous sensing and localized therapy.

Overall, conductive hydrogels integrate softness, high water content, adhesion and mixed ionic-electronic transport, making them especially well-suited for electrochemical interfaces with the brain, heart and skeletal muscle, where intimate, low-modulus contact to wet, excitable tissue is essential^[121]. As these materials evolve toward printable, adaptive and biodegradable architectures, they are expected to form a core platform for the next generation of implantable bioelectronics and soft robotic systems and to enable broader clinical and engineering applications^[122].

Liquid metals and their soft composites form a class of conductors that combine metal-level electrical conductivity with fluid- or rubber-like mechanics^[123]. In a narrow sense, liquid metals refer to metals or alloys that remain liquid near room temperature, most prominently gallium-based eutectic alloys such as eutectic gallium–indium (EGaIn) and Galinstan^[124,125]. These alloys exhibit conductivities on the order of 10⁶ S·m^-1 together with low vapour pressure, relatively low toxicity and good chemical stability, making them attractive as soft yet highly conductive layers in deformable bioelectronic systems.

Their functional behavior is governed by the interplay between metallic bonding and interfacial mechanics. Intrinsically, liquid metals are highly conductive fluids whose conductivity remains nearly unchanged under large mechanical deformation, allowing stable signal transmission under strain^[126]. A nanometre-scale oxide skin and high surface tension stabilise discrete droplets and filaments and allow damaged traces to self-heal as the liquid flows and closes cracks. In soft metal composites, micron- or nanoscale droplets dispersed in PDMS, PU or hydrogels remain electrically insulated at rest; under stretching, compression or shear, the oxide shells rupture and neighbouring droplets fuse into percolating networks, producing stress-activated conductivity and inherent strain sensitivity. These features are particularly beneficial at organ interfaces subjected to continuous motion, such as skin, skeletal muscle and cardiovascular tissues, where conductive layers must tolerate repeated deformation without loss of conductivity^[127]. For example, Zeng et al. showed that direct-ink-written EGaIn circuits can serve as highly stretchable electrodes in topology-optimized piezoelectric sensors for anisotropic motion monitoring, illustrating how liquid-metal interconnects enable precise patterning, stable conductivity under large deformation, and wearable bioelectronic integration [Figure 4C]^[128]. Beyond serving as passive fillers, gallium-based droplets can also initiate free-radical polymerisation and participate in crosslinking, as shown by Jaseem et al., who obtained tough, injectable and self-healing conductive hydrogels suitable for injectable electrodes and soft encapsulation layers on skin or muscle^[129].

Building on these platforms, recent work has shifted from simply achieving conductivity and stretchability to ensuring stable operation at complex biological interfaces. A notable example is a magnetically reshapable three-dimensional liquid-metal multi-electrode array, in which EGaIn-filled deformable microtubes are folded into three-dimensional shapes under magnetic guidance and gently inserted into brain organoids for multichannel, deep electrophysiological recording and stimulation, highlighting the plasticity and compliance of liquid metals at three-dimensional brain and organoid interfaces^[130]. Such demonstrations highlight the capacity of liquid metals to conform to soft, three-dimensional brain and heart while maintaining metallic-level conductivity, a property difficult to achieve with conventional rigid electrodes^[131,132].

However, liquid-metal conductors also present important fabrication and stability challenges. Their high surface tension and fluidity make high-resolution and reproducible patterning difficult, while the spontaneously formed gallium oxide skin can both stabilize traces and alter wettability, adhesion, and contact resistance^[133]. Poor wetting on polymer substrates may cause line retraction or discontinuous patterns, and leakage from microchannels or composite matrices under repeated deformation can compromise device reliability^[134]. These issues become more pronounced in large-area or high-throughput fabrication, where ink rheology, substrate adhesion, encapsulation, and droplet stabilization must be precisely controlled^[135]. Recent patterning and stabilization strategies, including surface modification, microfluidic confinement, transfer printing, direct ink writing, laser-assisted processing, and elastomer/hydrogel encapsulation, are therefore important for translating liquid-metal conductors into reliable organ-specific bioelectronic interfaces^[136].

Overall, liquid metals and their soft composites combine metal-like conductivity, rubber-like softness and reconfigurable interfaces, providing a conductive-layer strategy centered on strain-invariant conductivity and mechanical robustness in highly deformable biological systems^[137]. Within organ-specific bioelectronic chips, they are particularly suitable for interfaces that experience large or repetitive deformation, including epidermal and muscular electromyography (EMG) systems, cardiovascular monitoring devices, and emerging three-dimensional neural or organoid interfaces^[138,139]. As encapsulation strategies, droplet stabilization methods, and biocompatible formulations continue to improve, liquid-metal-based conductors are expected to expand from wearable platforms toward more demanding implantable organ interfaces^[140,141].

Two-dimensional conductors are atomically thin or few-layer materials in which charge carriers are largely confined to in-plane transport^[142,143]. Representative systems include graphene and its derivatives, transition metal dichalcogenides (TMDs) and MXenes^[144]. Compared with bulk conductors, these sheets offer very high specific surface area, tunable band structures and excellent flexibility: monolayer graphene combines high carrier mobility with nanometre-scale thickness, while MXenes provide hydrophilicity and easy dispersion through surface –O/–OH/–F terminations^[145]. These characteristics enable 2D materials to act as ultrathin conductive layers that enhance interfacial conductivity and device density without significantly increasing the bending stiffness of soft, organ-facing substrates.

Their interfacial behaviour is governed by sp²-conjugated carbon networks or transition-metal d orbitals together with engineerable defects and surface groups^[146,147]. Extended π or d bands support fast in-plane electronic transport, whereas edge sites, vacancies and terminations can be chemically tuned to adjust carrier density, Fermi level and charge-transfer kinetics. Reported conductivities vary widely depending on material type, flake quality, oxidation state, and film assembly; graphene-based and MXene films can reach conductivities from 10³ to 10⁵ S·m^-1 or higher in optimized films, whereas oxidized or defect-rich derivatives show substantially lower values^[19,148]. Unlike OMIECs, these materials do not usually exhibit bulk ionic–electronic mixed conduction, but their surface groups and electrolyte-accessible interfaces facilitate charge transfer and reduce electrode–tissue impedance^[149]. For example, Ding et al. integrated MXene nanosheets into a flexible multilayer sensing architecture composed of hydrogel microneedles, PET/ITO films, and PDMS encapsulation, where the layered MXene morphology, stress-induced interlayer transport, and non-faradaic interfacial coupling enabled deformable multimodal bioelectronic sensing [Figure 4D]^[150]. When coated onto elastomers or hydrogels, 2D flakes form composite films that markedly reduce electrode–tissue impedance and maintain stable conduction under repeated bending and stretching, supporting long-term mechanical compatibility at dynamic tissue interfaces^[104,151].

Beyond conductive 2D sheets, emerging two-dimensional polymers provide a distinct materials strategy for ultrathin encapsulation. For example, Ritt et al. showed that the molecularly impermeable 2D polyaramid 2DPA-1 can be processed into nanometre-thin, electrically insulating barrier films with exceptionally low gas permeability. A 60-nm 2DPA-1 coating markedly retarded MAPbI3 degradation by suppressing O₂ and water-vapour permeation, highlighting the promise of such ultrathin 2D polymer membranes for conformal encapsulation^[152]. This combination of conductive 2D layers with atomically thin barrier films offers a strategy for simultaneously improving electrical performance and protecting organ-resident devices from biofluid infiltration and molecular diffusion^[153].

Overall, two-dimensional conductors combine atomic-scale thickness, high conductivity and tunable interfacial chemistry, providing a conductive-layer strategy centred on high spatial resolution, minimal mechanical footprint and chemically programmable interfaces^[154,155]. In the brain and retina they can function as transparent, low-impedance electrodes and protective layers for high-fidelity recording and stimulation, whereas on the heart, skeletal muscle and intestinal mucosa, 2D conductor-elastomer or -hydrogel coatings can enhance electrical coupling and stability while minimally perturbing tissue mechanics^[156]. Within organ-specific bioelectronic chips, 2D materials are particularly advantageous when high-density electrode integration and minimal structural intrusion are required, such as in retinal prostheses, cortical microarrays and conformal visceral patches^[157].

Beyond mechanical and electrical matching, biocompatibility is a central design consideration in organ-specific bioelectronics^[158,159]. This issue extends beyond acute cytotoxicity to include skin compatibility, foreign-body and immune responses, inflammatory remodeling, and long-term safety during chronic use^[160]. These concerns depend strongly on the target organ, implantation mode, exposure duration, and local biochemical environment. Therefore, material selection should be evaluated not only by conductivity and mechanics, but also by the biological responses expected in each tissue context.

In addition to organ-specific constraints, bioelectronic materials must also be selected according to the intended modality. Recording, stimulation, biosensing, pressure sensing, temperature sensing, and delivery each impose distinct requirements for contact, impedance, charge injection, selectivity, stability, and transport. These modality-specific considerations are summarized in Table 2.

Photolithography is a widely used microfabrication technique that provides excellent spatial resolution by using patterned light to transfer features onto thin-film substrates^[184]. It is commonly applied to define metal electrodes and interconnects on polymeric substrates such as PI or SU-8, yielding flexible microelectrode arrays^[185,186]. This strategy has been used to fabricate high-density arrays for neural probes, retinal implants, and other interfaces, with feature sizes on the order of single cells and finely controlled electrode spacing, allowing highly localized interaction with biological tissues^[187]. A key limitation, however, is that conventional photolithography is largely restricted to planar, wafer-based processes; additional steps such as transfer printing or lamination are often required to integrate these micro-patterned structures onto organ-mimicking 3D surfaces.

Direct photopatterning offers a complementary route in which functional materials are structured in situ on a substrate without separate photoresist processing and etching [Figure 5A]^[188]. In this approach, light-sensitive formulations, such as photopolymerizable hydrogels containing conductive fillers or photo-crosslinkable conductive polymers, which are exposed through a mask or projected pattern to define the desired microstructures^[189]. This method allows conductive pathways, electrode sites, or microelectrode arrays to be “written” directly into soft, biocompatible matrices with high precision. For organ-specific bioelectronics, direct photopatterning is particularly attractive because it can produce very soft, tissue-like electrode architectures that conform to brain, cardiac, or other organ surfaces while minimizing mechanical mismatch, thereby improving the stability and fidelity of the tissue-electrode interface.

Laser writing (or laser engraving) employs focused laser beams to locally ablate, modify, or convert materials, thereby defining conductive tracks or electrode patterns without the need for masks. It is an inherently maskless and programmable technique that lends itself well to rapid prototyping of complex layouts. A widely studied example is laser-induced graphene (LIG), in which irradiation of polymeric or biomass-derived films converts the surface into a porous, conductive graphene network that can serve as a flexible electrode^[190]. This direct-write process can achieve microscale features and can be readily reconfigured at the design stage, which is advantageous for organ-specific devices that require iterative optimization or patient-specific tailoring [Figure 5B]^[191]. Moreover, laser writing can be performed on thin films that are later laminated onto curved tissues, or directly on preformed substrates intended to conform to skin or organ surfaces^[192]. In this way, it helps bridge the gap between high-resolution patterning and practical fabrication of conformal bioelectronic interfaces.

Inkjet printing is a digital, additive method where conductive or functional inks are deposited in tiny droplets to form patterns. Modern inkjet systems can produce microscale features, offering fairly high printing resolution for flexible circuits. This technique allows custom electrode layouts to be printed on demand, accommodating organ-specific geometries by simply adjusting the digital design. Because it is a non-contact process, inkjet printing can pattern delicate substrates without damage, making it suitable for creating fine electrode arrays on soft, tissue-like materials [Figure 5C]^[190]. Its capability to deposit multiple ink materials also enables integration of sensors or stimulators tailored to organ needs. For example, printing biocompatible conductive inks for cardiac or neural patches.

Electrospinning produces nonwoven mats of micro- or nanofibers by ejecting a polymer solution (often containing conductive nanomaterials) under a high-voltage field. The collected fiber mesh can cover large areas and serves as an inherently porous, flexible substrate or conductor network. Electrospinning is a low-cost, high-throughput technique that can be scaled up for mass production^[193]. The resulting fibrous networks have high surface area and can be made from biocompatible polymers, making them well-suited for interfaces with tissues and organs [Figure 5D]^[194]. In organ-specific bioelectronics, electrospun scaffolds can function as soft electrodes or supporting matrices that conform over an organ’s surface. For example, a cardiac patch with electrospun conductive fibers that wrap around the heart [Figure 5E]^[195]. Because the fiber diameter and mesh architecture can be tuned, this method allows optimization of mechanical properties (matching the compliance of an organ) while covering the target area uniformly with conductive pathways. Moreover, electrospun electronics often remain breathable and permeable, an advantage for long-term tissue integration.

“Cut-and-paste” fabrication is a paper-crafting-inspired approach where electronic patterns are first cut out of thin conductive films or foils and then transferred onto target substrates^[51,196]. In practice, this often uses a desktop vinyl cutter or laser cutter to outline circuits from metal-coated polymer sheets (such as gold on polymer). The patterned pieces (or the negative stencil) are then pasted onto surfaces like medical tape, skin patches, or organ models. This method bypasses the complexities of photolithography, achieving circuit feature sizes on the order of ~100 μm with simple equipment. It is highly suitable for large-area flexible electronics because it is not constrained by wafer size, allowing meter-long flexible circuits to be fabricated using roll-to-roll feedstocks^[183,197]. In organ-specific device fabrication, cut-and-paste approaches have been used to rapidly prototype epidermal sensors and electronic tattoos that conform to the skin, as demonstrated in epidermal electronic systems (EES) and related wearable devices^[198]. The same concept can also be extended to creating custom-shaped electronics for organs, such as cutting sensor meshes that match the geometry of heart or brain surfaces^[199]. The main advantages of this approach are speed and low cost, although the achievable resolution is lower than that of photolithographic methods. In addition, this technique can be integrated with other large-area fabrication strategies, for instance by using a cut stencil to spray-coat or print conductive inks in subsequent steps, thereby combining pattern definition with large-area surface coverage.

Chemical solution deposition involves coating a substrate with functional material from a liquid phase over a large area^[200]. Techniques in this class include spin coating, dip coating, blade coating (also known as doctor blading), and similar solution-based film deposition methods. They enable uniform, thin films of conductive or active materials (polymers, nanoparticle inks, etc.) to be laid down over device areas far larger than typical wafers. For instance, one can spin-coat a conductive polymer across an entire flexible sheet, or dip-coat a 3D object to blanket it with a conformal conductive layer. These processes are simple, rapid, and cost-effective, often requiring only basic lab equipment. In organ-bioelectronics, solution deposition can be used to create continuous electrodes or sensing skins that cover whole organ surfaces (like a “electronic membrane” for an organ). While the coatings themselves are unpatterned, they can be combined with shadow masks or post-deposition patterning (laser ablation, selective etching) to define electrode regions. An example application is coating a balloon catheter or a neural probe with an even layer of conductive polymer to improve electrical interface over its entire surface area.

3D printing (additive manufacturing) provides a direct route to build 3D device architectures that follow organ geometry. For organ-specific bioelectronics, this enables anatomically contoured implants, or electrode lattices whose overall shape and local electrode placement are defined in a single fabrication sequence. High-resolution methods such as two-photon polymerization (2PP) are particularly relevant because they produce sub-micrometer features and truly 3D structures on top of planar microelectronics. In a recent example, 2PP was used to print polymer microelectrode templates directly onto a CMOS array, followed by conformal metal coating, patterning, and passivation to yield 6,600 tissue-penetrating microelectrodes at 35 μm pitch, with independently tunable height and tip geometry^[204]. These pillar electrodes penetrate the retina and place the recording sites within the retinal ganglion cell layer while avoiding the axon bundle layer, enabling high-density, large-area recordings with reduced axonal interference. This type of “direct-print” 3D electrode strategy shows that additive manufacturing can separate the design of tissue-interfacing microstructures from the underlying electronics [Figure 6A]^[205]. It enables organ-matched curvature, depth targeting, and heterogeneous electrode shapes within a single array. These capabilities are difficult to achieve with planar lithography alone. For instance, researchers have printed flexible electronics in a dome shape to fit a cardiac ventricle^[201]. In this work, a scaled anatomical heart model is used as a 3D template for device construction. The electronic components are fabricated as serpentine metal meshes by conventional planar processing, but the overall 3D geometry and mechanical properties are set by a soft elastomeric membrane cast onto the 3D-printed heart model. After demolding, the electronics-integrated membrane can conformally wrap the entire epicardial surface, forming a stable bioelectronic interface with the heart. Beyond ex situ fabrication, recent advances have begun to extend additive manufacturing directly into living tissues. For example, Wei Gao and colleagues recently demonstrated in vivo ultrasound printing, where acoustic fields are used to localize and polymerize bioinks within deep tissues, enabling the formation of functional structures directly inside the body^[206]. This approach represents a paradigm shift from pre-fabricated implants toward minimally invasive, in situ fabrication of bioelectronic components, potentially allowing device formation in anatomically constrained or otherwise inaccessible regions [Figure 6B]. By choosing biocompatible, soft materials (e.g., stretchable inks or polymer composites), printed electronics can combine precise micro-patterning with bulk 3D shapes^[207]. In this way, 3D printing directly contributes to organ-interfacing devices by co-defining both device geometry and microstructure, reducing the need to bend or mechanically deform flat electronics to fit complex organ topographies^[208].

Spray coating (or spray printing) atomizes a solution or dispersion of functional materials into fine droplets and deposits them onto a target surface. Because the spray can wrap around edges and recessed regions, this method is well suited for coating substrates with irregular or curved geometries. In soft bioelectronics, spray processes are commonly used to form thin, percolating films of conductive nanomaterials [e.g., silver nanowires (AgNWs), graphene, carbon nanotubes] or polymer blends over large areas in a single step, while allowing control over thickness and sheet resistance through the number of passes and ink concentration^[209]. The technique is fast and can produce ultra-thin, uniform layers across the entire target, from convex domes to concave wells^[210]. Xu et al. exemplified this strategy by spray printing AgNWs onto multiscale porous styrene–ethylene–butylene–styrene (SEBS) elastomer substrates to construct on-skin electronic devices with integrated passive-cooling capability^[211]. The porous SEBS provided strong mechanical interlocking and π-π interactions with the AgNWs, allowing the sprayed networks to adhere firmly to the sponge-like surface while preserving high breathability and low thermal resistance. In this way, a single spray-coating step defined conductive traces and electrodes across complex, gas-permeable elastomer sheets, enabling multifunctional e-skin devices for electrophysiological sensing, temperature monitoring, and Joule heating without resorting to multi-step lithography. For organ-specific bioelectronics, spray coating can be applied to preformed 3D supports such as balloon catheters or molded elastomer shells^[212]. A conformal conductive or sensing layer can be deposited on the entire surface or in selected regions through patterned masks, while the underlying porous or compliant substrate sets the mechanical properties and curvature^[183,213]. Because spray processes are compatible with many polymers, hydrogels, and elastomers and are already mature in industrial settings^[214]. They provide a practical route to scalable, conformal coatings for implantable or on-organ devices that must match complex anatomy yet remain thin, breathable, and mechanically compliant^[203].

Conformal molding refers to fabrication schemes in which electronic structures are formed or transferred while already constrained by a 3D target geometry. Instead of fabricating devices only on flat wafers and later forcing them to bend, the substrate or stamp is shaped to match the desired curvature, and micropatterns are created or printed in this state. For example, Choi et al. fabricated electrodes, interconnects, and LEDs on thin elastomeric films and then thermoformed the patterned circuits onto curved 3D molds, enabling conformal electronic systems on complex curvilinear surfaces [Figure 6C]^[215]. Alternative approaches use compliant stamps bearing micro- and nanopatterns (for example, PDMS or other elastomers); under heat and pressure, these stamps conform to spherical or highly curved surfaces and imprint conductive features directly onto them^[184]. In all cases, the goal is to obtain intimate, continuous contact over the full 3D surface, which is essential for stable electrical coupling and reduction of local stress concentrations at the tissue-device interface.

Conformal additive stamp (CAS) printing provides a representative example of this concept applied with wafer-level precision. In CAS printing, ultrathin devices or “inks” are first fabricated on planar donor wafers using standard microfabrication, then retrieved by a pneumatically inflated elastomeric balloon and printed onto non-developable 3D substrates such as hemispherical shells and contact-lens molds. The balloon stamp deforms to match complex curvilinear surfaces while keeping the strain in brittle silicon elements below fracture limits, enabling high-yield transfer of photodetector arrays, antennas, hemispherical solar cells, and multifunctional smart contact lenses^[216,217]. For organ-specific bioelectronics, similar conformal molding strategies allow sensor grids and electrode meshes to be fabricated in pre-curved forms that match the epicardium, ocular globe, or visceral organs^{[201,216,218]}. Once released from the mold, these meshes “hug” the organ with minimal additional fixation and can be combined with stretchable designs to accommodate physiological motion while maintaining the imposed 3D shape.