Introduction

The integration of electronics and biological systems has driven remarkable advancements in bioelectronics, enabling the monitoring, regulation, and interaction of living organisms through electronic interfaces1,2,3,4. The earliest reported bioelectronic experiment was Galvani’s demonstration of generating frog leg movements through electric stimulation5, which revealed the interplay between physiological activity and electricity. Today, bioelectronics has evolved into a multidisciplinary domain with significant academic and commercial impacts. For example, wearable sensors facilitate continuous body signal collection in daily life with minimal interference6,7,8, offering transformative potential in daily healthcare9,10, human–machine interfaces11,12,13, and augmented/virtual reality14,15. Recent research has further extended bioelectronic applications to various animals16,17,18, insects19,20, plants21,22,23, and fungi24,25, enriching our toolbox for interacting with the living world.

The geometric mismatch between biosurfaces and conventional electronics poses a significant challenge for the integration of bioelectronics7,26,27. Biosurfaces vary widely in size, curvature, roughness, and microstructure, whereas traditional electronics are typically rigid and bulky26,27. Fixing the conventional electronics to curved biosurfaces requires the use of adhesive tape, fixing garments, or suturing28, which leads to an increase in total size/weight, user discomfort, or even an inflammatory response. On rough or textured surfaces, such as human skin, rigid devices tend to form interfacial gaps that compromise device performance12,29. Moreover, the dynamic and continuous deformation of biological tissues introduces motion artefacts and reduces the long-term stability of non-conformal bioelectronic systems30,31. Emerging soft electronics offer a promising solution for stable integration by deformation to match the contours of biological tissues2,32,33. Enhanced conformability also improves the performance of bioelectronics. For instance, conformable soft electrodes establish stable electrical interfaces with the skin or organ surface, facilitating high-fidelity recording of electrophysiological signals during physiological activities29,34,35,36,37. Compared with conventional electrodes30,31, conformal soft electrodes avoid the use of additional fixation components, such as adhesive tape, conductive paste, or conductive gel, which enables long-term stability, low motion artefacts, low inflammation risk, and improved user comfort.

Novel materials, geometric designs, and technologies have been developed for conformal bioelectronics. In this review, we focus on the strategies for improving the conformability of bioelectronics to complex biosurfaces (Fig. 1). First, theoretical models are discussed to understand the theoretical basis of conformable bioelectronics. Subsequently, we introduce common strategies for achieving conformability, including softness improvement, interfacial adhesion enhancement, and Gaussian curvature mismatch reduction. Finally, we summarise and discuss the prospects of conformable bioelectronics.

Fig. 1: Conforming strategies for bioelectronics on arbitrary surfaces.
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Reproduced with permission from ref. 12. Copyright 2013 Wiley-VCH. Reproduced with permission from ref. 78. Copyright 2015 Springer Nature. Reproduced with permission from ref. 23. Copyright 2018 Springer Nature. Reproduced with permission from ref. 212. Copyright 2023 Wiley-VCH. Reproduced with permission from ref. 227. Copyright 2021 Wiley-VCH. Reproduced with permission from ref. 183. Copyright 2016 American Chemical Society. Reproduced with permission from ref. 260. Copyright 2021 Springer Nature. Reproduced with permission from ref. 263. Copyright 2021 Wiley-VCH. Reproduced with permission from ref. 48. Copyright 2014 American Society of Mechanical Engineers. Reproduced with permission from ref. 54. Copyright 2008 Royal Society of London. Reproduced with permission from ref. 119. Copyright 2020 American Association for the Advancement of Science. Reproduced with permission from ref. 264. Copyright 2023 American Association for the Advancement of Science. Reproduced with permission from ref. 276. Copyright 2024 Springer Nature.

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Theoretical model of conformability

Biosurfaces have complicated geometric features that hinder seamless bioelectronic integration. For example, biosurfaces span a wide range of curvature radii, from millimetre-scale neural fibres36 (Fig. 2a) to centimetre-scale organs34,35,38 (Fig. 2b), to the relatively flat surfaces of the chest or back39,40 (Fig. 2c), resulting in variations in the required bendability of bioelectronic devices to achieve optimal conformability. Additionally, biosurfaces possess different surface textures, which influence the conformal attachment of bioelectronics. For instance, the surfaces of nerves and blood vessels are smooth41, whereas the surface of the skin is often rough with an amplitude of 15–100 μm 2,42 (Fig. 2d). The microstructure of the biosurface further introduces complexity for conformal attachment, such as sweat pores43 (Fig. 2e), leaf stoma21 (Fig. 2f), and hair13,44,45,46. Unlike curved biosurfaces, most soft electronics are currently fabricated on flat substrates for manufacturing simplicity. Such a geometric mismatch results in the necessity of shape transformation when bioelectronics are attached to the biosurface26,47, which is governed by various factors, such as substrate geometry, mechanical properties of bioelectronics, and the interaction between bioelectronics and biosurfaces.

Fig. 2: Examples of conformal bioelectronics on biosurfaces showing various geometry features.
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a A neural electrode twisted on millimetre-sized neural fibre. Reproduced with permission from ref. 303. Copyright 2019 American Association for the Advancement of Science. b Smart wristband mounted with bending radius of several centimetres. Reproduced with permission from ref. 304. Copyright 2016 Springer Nature. c Flexible electrocardiogram electrodes on nearly flat chest surfaces. Reproduced with permission from ref. 39. Copyright 2023 Wiley-VCH. d Flexible mesh conformal to rough skin surface. Reproduced with permission from ref. 270. Copyright 2013 Wiley-VCH. e Gold nanomesh conformal to fingerprint (left) and sweat pore on the fingerprint (right). Reproduced with permission from ref. 43. Copyright 2017 Springer Nature. f Silver nanowire electrodes that covers the stoma on the leaf. Reproduced with permission from ref. 21. Copyright 2025 Springer Nature.

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To systematically understand and optimise conformability, researchers have developed various theoretical models to analyse the bioelectronic–biosurface systems. Most theoretical models have been simplified to focus on the major features of the target system. These models provide valuable design guidelines for creating bioelectronics that can seamlessly adapt to biosurfaces while maintaining their functionality. In the following section, we review key theoretical frameworks that have advanced our understanding of conformable electronics and their biological integration.

Conformability on rough surfaces

To analyse the conformability of bioelectronics on rough biosurfaces such as the skin, researchers have developed two-dimensional theoretical models. Wang et al. proposed a skin model with a sinusoidal profile described by \(y=(1+\cos (2\pi x/\lambda ))h/2\) to represent surface roughness (Fig. 3a, b), where h represents the wrinkle amplitude, and λ represents the wavelength48,49. In this model, thin-film bioelectronics are simplified as elastic beams that perfectly conform to the skin surface after lamination. The total energy of the complete conformal system is expressed as \({\bar{U}}_{\text{conformal}}={\bar{U}}_{\text{bending}}+{\bar{U}}_{\text{skin}}+{\bar{U}}_{\text{adhesion}}\), where \({\bar{U}}_{\text{bending}}\), \({\bar{U}}_{\text{skin}}\), and \({\bar{U}}_{\text{adhesion}}\) represent the bending energy of the bioelectronics, elastic energy of the skin, and interfacial adhesion energy between the skin and bioelectronics, respectively. Given the conformal criterion of \({\bar{U}}_{\text{conformal}} < 0\), the above model derives the conformable criterion for bioelectronics \(\frac{\pi {h}^{2}}{\gamma \lambda } < \frac{16}{{E}_{\text{skin}}}+\frac{{\lambda }^{3}}{{\pi }^{3}{EI}\,}\), in which \(\gamma\), \({E}_{\text{skin}}\), and \({EI}\) represent the skin-electronics interfacial energy coefficient, Young’s modulus of the skin, and the effective bending stiffness of the bioelectronics, respectively.

Fig. 3: Theoretical models for conformal bioelectronics.
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a Schematic illustration showing a thin film fully conformal to a wavy skin surface. Reproduced with permission from ref. 48. Copyright 2014 American Society of Mechanical Engineers. b Theoretical predicted conformability of polyimide-gold epidermal electronics on skin. Reproduced with permission from ref. 48. Copyright 2014 American Society of Mechanical Engineers. c Schematic illustration showing a thin film partially conformal to a wavy skin surface50. Reproduced with permission from ref. 50. Copyright 2016 American Society of Mechanical Engineers. d Theoretical predicted conformability on a typical skin surface. Reproduced with permission from ref. 50. Copyright 2016 American Society of Mechanical Engineers. e Schematic illustration showing a thin film on a sphere surface. Reproduced with permission from ref. 54. Copyright 2008 the Royal Society (U.K.). f Experimental result of conformal sizes of thin films attached to a sphere surface. Reproduced with permission from ref. 59. Copyright 2011 American Physical Society. g Experimental and simulation results of thin film and edge-cut thin film attached to a sphere surface. Reproduced with permission from ref. 53. Copyright 2023 American Association for the Advancement of Science.

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Subsequent studies have developed more sophisticated models to enhance accuracy and broaden applicability50,51,52. For example, Wang and Lu considered the elastic energy of bioelectronics and partially conformed cases50 (Fig. 3c), revealing the conformability of a given system is controlled by four dimensionless parameters (Fig. 3d): the wrinkle aspect ratio \(\beta =2\pi h/\lambda\), thickness–wavelength ratio \(\eta =t/\lambda\), modulus ratio \(\alpha ={E}_{m}/{E}_{s}\), and adhesion parameter \(\mu =\gamma /({E}_{s}\lambda )\). In addition, the contact zone size \(\hat{x}=2{x}_{c}\lambda\) can be derived through energy minimisation of the model. Cai et al. leveraged a moderately large deflection model for bioelectronics to improve model accuracy in ultrathin bioelectronic systems52, showing that such a system is determined by the normalised adhesion energy \(\bar{\gamma }=(24\gamma {\lambda }^{4})/(\pi \bar{E}{h}^{3}{A}^{2})\), conformability ratio \(b/\lambda\), and thickness–amplitude ratio \(h/A\).

Global conformability on non-developable surfaces

Biosurfaces can be categorised into developable and non-developable surfaces according to their geometric properties. Developable surfaces, such as plane and cylindrical surfaces, have zero Gaussian curvature at every point on the surface. In contrast, non-developable surfaces, such as spherical and saddle surfaces, have non-zero Gaussian curvatures. On the other hand, bioelectronics often feature a developable thin-film geometry, as they are initially fabricated on a flat substrate. Because such bioelectronics are generally bendable but non-extensible, mounting bioelectronics on biosurfaces can be considered an isometric transformation process. According to Gauss’s Theorema Egregium, such a transformation maintains the local Gaussian curvature, i.e., the developable thin-film bioelectronics can be conformally attached only to developable biosurfaces without any in-plane compressing or stretching deformation47,53. However, practical implementations often overcome this geometric constraint through material compliance. Both bioelectronic devices and biological tissues typically possess a certain stretchability to accommodate local deformations, thereby enabling conformal attachment, even when geometric mismatches exist between developable devices and non-developable biosurfaces54,55,56.

Several models have been developed to analyse the attachment of thin-film devices to non-developable surfaces53,54,57,58. For example, Majidi and Fearing54 analysed the mounting of a circular thin film on a rigid sphere with radii of \({R}_{f}\) and \({R}_{s}\), respectively (Fig. 3e). By applying the von Kármán plate model to the thin film that has a thickness of \(h\), Young’s modulus of \(E\), Poisson’s ratio of \(\nu\), and interfacial energy coefficient of \(\lambda\), we can derive the stable criterion condition for full conformability \(\frac{{R}_{\text{f}}^{4}}{128{R}_{\text{s}}^{4}}+\frac{{h}^{2}}{12\left(1-\nu \right){R}_{\text{s}}^{2}}\le \frac{\lambda }{{Eh}}\). Therefore, to achieve an optimal conformal attachment to a spherical surface, a small size ratio \({R}_{f}/{R}_{s}\), small thickness h, and soft material with a low modulus E are preferred for thin-film bioelectronics. For thin films with a negligible thickness, the above model provides a critical conformable radius ratio \({R}_{s}/{R}_{f}\) that scales with \({\left(\lambda /{Eh}\right)}^{\frac{1}{4}}\), which was verified by experiments59 (Fig. 3f). Meanwhile, for the system with minimised energy, the in-plane radical and hoop strain are \({\varepsilon }_{r}={\left(\frac{{R}_{{\rm{f}}}}{{R}_{{\rm{s}}}}\right)}^{2}\frac{\left(1-v\right)-\left(1-3v\right){\left(r/{R}_{{\rm{f}}}\right)}^{2}}{16}+\frac{z}{{R}_{{\rm{s}}}}\) and \({\varepsilon }_{{\rm{\theta }}}={\left(\frac{{R}_{{\rm{f}}}}{{R}_{{\rm{s}}}}\right)}^{2}\frac{\left(1-v\right)-\left(3-v\right){\left(r/{R}_{{\rm{f}}}\right)}^{2}}{16}+\frac{z}{{R}_{{\rm{s}}}}\) for points located at the radical coordinate of r and height coordinate of z, respectively54,56. These expressions highlight the significant strain development during the spherical mounting of thin-film devices, which is significantly influenced by the lateral size of the thin film device. Liu et al. studied a film-on-sphere system through experiments and coarse-grained molecular dynamics simulations53 (Fig. 3g), revealing that a critical maximum radius exists to achieve full conformability for a given system. A film with a larger radius delaminates from its film. Furthermore, such delamination can be reduced by cutting several radial slits on the film to release the stress53. Wang et al. developed a general model to analyse the attachment of a thin-film device to a substrate with arbitrary surfaces57, emphasising the general effect of a smaller lateral film size, film thickness, and larger interfacial interaction to achieve good conformability.

In summary, among the various theoretical models, the thickness, elasticity, and interfacial interaction with a biosurface determine the conformability of thin-film bioelectronics. Moreover, for planar bioelectronics attached to a non-developable surface, lateral geometry emerges as an additional critical parameter. In the following sections, we review the conformability improvement strategies, including softness improvement, interfacial adhesion enhancement, and Gaussian curvature mismatch reduction.

Softness improvement

When attached to a curved biosurface, bioelectronics frequently undergo shape transformation, which leads to the accumulation of bending and elastic energy. The deformational potential energy can be reduced by improving the bending and elastic softness of bioelectronics, which promotes the conformability.

Geometry design

Geometric design is a key factor of the mechanical softness in bioelectronics. For instance, the bending stiffness of a thin film device can be described as \(D=E{h}^{3}/12\left(1-{\nu }^{2}\right)\), indicating that a smaller thickness leads to higher bendability. Moreover, engineered geometries such as serpentine interconnects, mesh layouts, and other deformable patterns can further enhance compliance. This section reviews representative geometry design strategies for improving softness in bioelectronics. Table 1 summarises typical soft bioelectronics that employ these principles.

Table 1 Comparison of typical conformable bioelectronics with geometry design.
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Thickness reduction

Reducing the device thickness is one of the most effective strategies for improving bioelectronic conformability, as both bending and stretching stiffness scale dramatically with thickness26,60,61. For example, thin-film bioelectronics fabricated on a 30-μm polyester substrate exhibit excellent mechanical properties matching with the skin2 (Fig. 4a), enabling the first demonstration of epidermal electronics. Jeong et al. systematically investigated this relationship by developing silicone-based thin-film bioelectronics with varying thicknesses and quantifying their conformability to skin surfaces12. Experimental results show that 5-μm bioelectronics exhibit perfect conformability with the skin, whereas devices with a thickness of 36 μm or larger exhibit progressively larger gaps at the skin–device interface (Fig. 4b). Such results also agree well with the theoretical model, which indicates a critical thickness of 25 μm (Fig. 4c). Compared with non-conformable devices, conformable 5-μm devices achieve a reduction in contact impedance (from ~90 kΩ to 24.4 kΩ), which enabling collecting biosignals with a low background noise of 12 μVRMS. Yokota et al. demonstrated ultra-conformable multi-layered organic light-emitting diodes (OLEDs) using 1-μm parylene as substrates62, achieving a total device thickness of 3 μm (Fig. 4d). The ultrathin display maintains its functionality while seamlessly conforming to the complex topography of human skin, enabling novel applications in wearable displays.

Fig. 4: Softness improvement via reduction of thickness.
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a Epidermal electronics fabricated on a 30-μm polyester substrate. Reproduced with permission from ref. 2. Copyright 2011 American Association for the Advancement of Science. b Conformability of silicone rubber thin film with different thickness on a skin replica. Reproduced with permission from ref. 12. Copyright 2013 Wiley-VCH. c Theoretical prediction of conformability in (b). Reproduced with permission from ref. 12. Copyright 2013 Wiley-VCH. d Ultrathin optoelectronics fabricated on 1-μm parylene substrate. Reproduced with permission from ref. 62. Copyright 2016 American Association for the Advancement of Science. e Conformability of cortical electrode array with different substrate thickness. Reproduced with permission from ref. 16. Copyright 2010 Springer Nature. f Mesh electrode array conformably attached on brain surface and collected neural activity signals. Reproduced with permission from ref. 16. Copyright 2010 Springer Nature.

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Cancelling the substrate of thin-film electronics further reduces the thickness of bioelectronics7,63, thereby improving conformability. This strategy typically employs sacrificial substrate layers that provide temporary mechanical support during device fabrication and handling and are subsequently removed after deployment16,43,64. For example, biocompatible, flexible, and solution-processable silk thin films have been used as temporary substrates for large-area cortical electrode arrays16. After device deployment on the brain surface, the silk substrate dissolves upon exposure to warm water, leaving a substrate-free electrode array that achieves superior conformal contact with the highly curved cortical surface (Fig. 4e). Compared with electrode arrays associated with a polyimide thin films, the substrate-less electrode array exhibits significantly improved conformability, therefore offering better data quality with high mean rms amplitude of 5.7 ± 3.0 across nearly all the channels (Fig. 4f). Alternatively, functional devices can be fabricated directly on target surfaces, eliminating the necessity for supporting substrates. For example, silver nanowires–polyurethane ink can be spray-coated onto human skin to fabricate a substrateless sensor65 that collects essential strain signals for hand gesture recognition tasks.

Deformable patterns

The incorporation of a patterned architecture is an effective approach to enhancing the softness of bioelectronic devices. Such engineered structures effectively reduce the overall stiffness while simultaneously enabling complex deformation modes that allow intrinsically brittle materials to achieve remarkable extensibility66,67,68. Among the various designs, serpentine patterns are frequently used in soft electronics69,70,71,72. For instance, serpentine patterned gold structure embedded in a polydimethylsiloxane (PDMS) matrix enables conductivity up to a 54% stretching69 (Fig. 5a), which is significantly higher than the inherent stretchability of unstructured gold thin films (<3%)73. Researchers have further enhanced the stretchability of serpentine designs using fractal structural motifs74,75,76,77 (Fig. 5b). Jang et al. designed a network structure based on a serpentine-shaped unit that enables omnidirectional stretchability78 (Fig. 5c). Networks with other deformable units, such as hexagonal units79, square units61,80,81,82, and hierarchical designs83,84, have also been incorporated into soft bioelectronics.

Fig. 5: Softness improvement via deformable pattern design.
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a Stretching of straight/serpentine gold trace. Reproduced with permission from ref. 69. Copyright 2004 Wiley-VCH. b Ultra stretchable fractal design based on serpentine unit. Reproduced with permission from ref. 77. Copyright 2013 Elsevier. Reproduced with permission from ref. 75. Copyright 2015 Springer Nature. c Omnidirectional stretchable network. Reproduced with permission from ref. 78. Copyright 2015 Springer Nature.

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Origami and kirigami structures have been adopted from papercraft fabrication to introduce controllable reconfigurability in thin-film devices85,86 and show potential for soft electronics. In origami structures, folding patterns are introduced into flat devices at predesigned locations to allow programmable deformation. For example, the Miura-ori design enables reversible folding of sheets into smaller areas with predictable pathways87. By leveraging the Miura-ori design, Hou et al. developed flexible thermal electric devices that achieve conformal attachment to a heat source with 3D curved surfaces, such as a human arm88. In contrast, kirigami structures involve the periodic cutting of thin films, which enhances stretchability86,89. Compared with other stretchable pattern designs, the kirigami structure offers a large effective area, which provides unique advantages for the surface electrophysiological interface90,91. By introducing a 2D fractal cutting design, biaxial extensibility can be achieved in thin-film devices92.

Textile electronics

Textile electronics is a promising platform for conformal bioelectronics6,9,93,94,95,96. Various textile-based electronics have been demonstrated, including conformal skin electrodes97,98, sensors99,100, energy-management devices101, displays102, and computing units103. The unique performance of textile electronics results from their hierarchical structures, which offer significant opportunities for achieving novel functions through special designs on multiple length scales. For example, electronic textiles with a knitted structure exhibit high stretchability because the knitted structure is composed of S-shaped fibre units97,104,105 (Fig. 6a). Based on the multi-level subsequent cracking of the conductive path (Fig. 6b), fabric strain sensors with woven structures exhibit high sensitivity over a wide range (≈500%)99,100. Moreover, by assembling functional fibres, various electrical devices, such as electroluminescent units102 (Fig. 6c), memristors106, pressure sensors107, and organic electrochemical transistors (OECTs)108,109, could be fabricated on the crossbar point110. Beyond traditional textiles that are commonly based on micrometre-sized fibres, electrospinning technology enables the fabrication of non-woven fabric composed of ultrathin fibres with diameters below 1 μm, offering a textile platform for conformal electronics with higher softness95,111,112,113,114. Conductive nanomesh has been developed by depositing conductive components on the nanofibres43,113,115, enabling the development of various body-conformal electronics, such as electrophysiological sensors43,116, hydration sensors64, strain sensors117,118, pressure sensors119, and OECTs120. In addition, carbon-based nanofibre textile electronics can be produced by converting non-conductive nanofibre precursors to conductive nanofibres via carbonisation121,122.

Fig. 6: Textile electronics.
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a Scanning electron microscopy image and schematic illustration of a weft knitted stretchable conductive textile. Reproduced with permission from ref. 105. Copyright 2017 Wiley-VCH. b The optical microscopy image, sensing mechanism, sensing performance, and pulse monitoring of a woven fabric sensor. Reproduced with permission from ref. 100. Copyright 2017 the Royal Society of Chemistry. c Schematic illustration and optical image of a textile electro luminescent display. Reproduced with permission from ref. 102. Copyright 2021 Springer Nature. d Optical microscope image, and skin sensor application of a conductive nanomesh fabricated on polyurethane-polydimethylsiloxane nanomembrane. Reproduced with permission from ref. 117. Copyright 2020 American Association for the Advancement of Science.

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Flexible material

Conventional electronics frequently rely on inorganic materials, such as copper, gold, silver, and silicon, which offer excellent electrical properties and well-established fabrication techniques. However, these traditional materials suffer from intrinsic mechanical brittleness, which limits their application in conformal bioelectronics. For instance, the Young’s modulus of conventional electronic materials typically exceeds 10 GPa, whereas that of human soft tissue is generally lower than 100 kPa. The mismatch between conventional electronic materials and biotissues impedes the conformal mounting of bioelectronics. Additionally, materials with a high Young’s modulus are prone to induce foreign body responses, which hinders implanted bioelectronic applications. Recent advances have focused on intrinsically soft materials that combine electrical functionality with mechanical compliance3,123,124, such as conductive polymers125,126,127, nanomaterials128,129,130, hydrogels131,132, and liquid metals133,134. These innovative soft electronic materials offer exceptional potential for achieving soft conformable devices capable of seamless integration with biological systems. Table 2 summarises the typical materials for bioelectronics and their physical properties.

Table 2 Physical properties of typical electronic materials.
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Substrate material

Polyimide thin films are widely used in bioelectronics owing to their excellent stability, commercial availability, and well-established processing techniques56,135,136,137,138. However, their high Young’s modulus limits the applicability of polyimide-based devices in conformal bioelectronics. In contrast, PDMS has become a common substrate material for conformal applications because of its low Young’s modulus, controllable curing process, good stability, and excellent biocompatibility12,70,78,139,140,141. Other synthetic elastomers, such as polyurethane117,142 (PU), thermoplastic polyurethane143,144 (TPU), and styrene-ethylene-butylene-styrene145,146,147 (SEBS), have also demonstrated applications in conformal bioelectronics. In addition, bioelectronics based on biopolymer substrates have also demonstrated extraordinary biocompatibility and sustainability148, such as silk proteins16,149,150,151, cellulose152, and collagen153. Recently, reconfigurable materials have been used to improve the conformability of bioelectronics via shape morphing124,154,155,156,157,158. For instance, Yi et al. developed a water responsive super contractile film as a substrate for bioelectronics154, which enables the conformal wrapping of bioelectronics on target biological objects such as nerves, muscles, and the heart. Self-healing polymer substrates can further improve conformability by adapting to complex biosurfaces124,155,156,158. Moreover, the viscoplasticity of self-healing polymer effectively dissipates deformation-induced elastic energy, enabling stress-free conformal bioelectronics.

Conductive polymers

Conductive polymers possess unique electrical properties and polymer-like mechanical properties given their special conjugated polymer structure33,127, endowing them with potential in bioelectronics125. Many conductive polymer families have been developed126,159,160,161 (Fig. 7a), including polyacetylene, polyaniline (PAni), polypyrrole (PPy), and poly(3,4-ethylenedioxythiophene) (PEDOT). Notably, additives are frequently used to improve mechanical properties, electrical properties, and processability162,163. For example, pristine PEDOT is frequently doped with polystyrene sulfonate (PSS) to enable oxidisation-induced doping and stable water dispersion126, resulting in one of the most frequently used conductive polymers PEDOT:PSS. The introduction of waterborne polyurethane and D-sorbitol to PEDOT:PSS further improves its flexibility and adhesion to the skin, enabling ultrathin conformal dry electrodes capable of biosignal monitoring164 (Fig. 7b). Moreover, conductive polymers are frequently solution-processable, allowing cost-effective and scalable fabrication through versatile liquid-phase techniques such as spin coating151,160, blade coating165,166, inkjet printing167,168,169, and dip coating170. For instance, Li et al. achieved highly thermal-wet comfortable and conformal electrodes by coating silk nanofibre mats with a PEDOT:PSS solution170 (Fig. 7c). Cui et al. fabricated transparent electrodes based on spin-coated PEDOT:PSS and silk hydrogel thin films (Fig. 7d), which conform well to brain surfaces, enabling the recording of neural activities151. As the synthesis of conductive polymers involves considerable amounts of solvents, chemical waste, and high-cost reactions, the sustainability of conductive polymer-based bioelectronics has become an important challenge125,127. Park et al. developed a closed-loop recyclable flexible electronic device based on selected materials that can be recaptured and reused with eco-friendly solvents171, demonstrating various applications such as electrophysiological electrodes, physical sensors, transistors, and inverters.

Fig. 7: Soft electronics based on conductive polymers.
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a Typical structures of conductive polymers. b Self-adhesive dry electrode based on PEDOT:PSS. Reproduced with permission from ref. 164. Copyright 2020 Springer Nature. c Thermal-wet comfortable skin electrodes based on PEDOT:PSS and silk nanofibre. Reproduced with permission from ref. 170. Copyright 2021 American Chemical Society. d Ultrathin transparent electrode based on PEDOT:PSS and PEGylated silk fibroin. Reproduced with permission from ref. 151. Copyright 2013 Wiley-VCH.

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Nanomaterials

Nanomaterials are also attractive candidates for the functional part of bioelectronics123,172,173,174, such as metallic nanomaterials129,175,176, carbon-based nanomaterials128,177,178, and MXenes179. Given their reduced size, nanomaterials offer improved flexibility compared with the corresponding bulk materials123. Silver nanowires (AgNWs) are among the most frequently used nanomaterials in bioelectronics owing to their excellent electrical conductivity, flexibility, and feasibility for liquid-based processes129,175,180,181. Moreover, the high conductivity and 1D geometry of the AgNWs allow the formation of a sparse conductive network, which enables the fabrication of transparent flexible electrodes176,182. For example, Kang et al. fabricated a transparent electrode by embedding an orthogonal AgNW array in an ultrathin polymer matrix176, which exhibits a transparency over 90% (Fig. 8a). Given its small thickness of ~100 nm, the AgNW array electrode achieves conformal attachment to the skin surface, resulting in imperceptible acoustic devices, such as loud speakers and microphones. He et al. employed a substrate-less, transparent AgNW electrode to monitor plant without interfering with photosynthesis21 (Fig. 8b). One of the major challenges for AgNW-based bioelectronics is stability, as AgNWs are easily oxidised, particularly for applications associated with biofluids such as sweat. Recently, surface modifications such as biocompatibility, stability, and conductivity have been used to further improve the properties of AgNWs. Choi et al. developed an elastic conductor based on the combination of the Au–Ag sheath-core nanowires and elastomer23 (Fig. 8c). Compared with bare AgNWs, Au–Ag nanowires exhibit higher stability and biocompatibility, allowing applications in conformal skin electronics and implanted electronics.

Fig. 8: Soft electronics based on nanomaterials.
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a Transparent electrodes based on orthogonal silver nanowire arrays and its on-skin acoustic application. Reproduced with permission from ref. 176. Copyright 2018 American Association for the Advancement of Science. b Transparent and breathable Ag nanowire electrode for plant monitoring. Reproduced with permission from ref. 21. Copyright 2025 Springer Nature. c Au-Ag sheath-core nanowire for elastic skin sensor. Reproduced with permission from ref. 23. Copyright 2018 Springer Nature. d Graphene electronic tattoo sensor and its electrocardiography application. Reproduced with permission from ref. 187. Copyright 2017 American Chemical Society. e Ultrathin stretchable electronic interface based on laser induced graphene and hydrogel. Reproduced with permission from ref. 141. Copyright 2024 Springer Nature. f Adhesive underwater epidermal sensor based on graphene. Reproduced with permission from ref. 197. Copyright 2023 Wiley-VCH. g Left: Printable highly stretchable conductive composite based on silver flakes and elastomer. Reproduced with permission from ref. 202. Copyright 2017 Springer Nature. Right: Printed stretchable conductor. Reproduced with permission from ref. 302. Copyright 2015 Springer Nature. h Self-healable electronic tattoo based on silk-graphene composites. Reproduced with permission from ref. 190. Copyright 2019 Wiley-VCH.

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Carbon nanomaterials are another important category of active materials for bioelectronics owing to their combination of flexible and strong mechanical properties, tuneable electrical properties, low weight, and environmental stability128,177. Various carbon nanomaterials have been utilised in bioelectronics, such as carbon nanotubes33,98,183,184,185,186 (CNTs), graphene (including pristine graphene114,187,188,189, reduced graphene oxide (rGO)104,190, and laser-induced graphene (LIG)141,191,192,193), and biopolymer-derived nanocarbons194. Notably, given the ultrahigh Young’s moduli of carbon nanomaterials195,196, ultra-thin geometries are frequently adopted to improve their softness. For example, an ultra-flexible graphene electronic tattoo was fabricated by laminating monolayer graphene onto a sub-micrometre poly(methyl methacrylate) (PMMA) film11,187,188 (Fig. 8d), which successfully enabled conformal skin sensors for electrophysiological signal. Lu et al. laminated an LIG thin layer on ultrathin hydrogel substrate (~1 μm) by cryogenically transferring for stretchable interfaces141, which shows high adhesiveness, stretchability, and conductivity (Fig. 8e). Wang et al. sandwiched a thin graphene layer between the elastic silicone layers to fabricate an epidermal sensor197 (Fig. 8f). With the assistance of an adhesive layer, these sensors successfully achieve stable conformal attachment to the skin, enabling underwater sensing. Moreover, based on the arrangement of carbon atoms, carbon nanomaterials have different electrical properties spanning from conductors to semiconductors, enabling their wide applications in bioelectronics, such as electrodes33, transistors198,199, and sensors189.

Notably, nanomaterials are often blended with other additives to improve their processability, mechanical properties, and biocompatibility148,200,201. For example, the combination of conductive silver flakes, elastic fluoroelastomer, and surfactant leads to a stretchable conductive composite with an initial conductivity of 4000 S/cm and a conductivity of 935 S/cm when subjected to a 400% strain202 (Fig. 8g), which enables a stretchable sensor network on gloves. Moreover, the uniform and stable composite ink is compatible with the printing process, which is compatible with large-scale production. The hydrophobic surfaces of carbon nanomaterials can be decorated with biomolecules98,203,204,205, which enables stable aqueous ink with improved biocompatibility and sustainability. Such carbon nanomaterial ink has versatile applications in printed conformal bioelectronics such as skin electrodes190 (Fig. 8h), textile electronics97,98, and breath sensors98,206. Notably, the incorporation of recyclable polymer binders enables the recycling of expensive nanomaterials207,208, providing a possible solution for sustainable and economical soft bioelectronics.

Hydrogels

Hydrogels feature an ultra-low modulus owing to their special structure of 3D crosslinked polymer network and interstitial water, making them particularly suitable for ultra-conformable bioelectronic applications131,132,209,210,211. Recent advances have demonstrated the successful integration of thin-film optoelectronics with sub-10 μm hydrogel substrates212 (Fig. 9a), significantly improving both device adhesion to skin and overall biocompatibility. The functional versatility of hydrogels can be further enhanced by incorporating conductive fillers, including conductive polymers213 (Fig. 9b), nanomaterials214 (Fig. 9c), or ionic salts112,211,215, to create conductive hydrogel composites. Additionally, the introduction of dynamic bonds within the polymer network enables the development of self-healing bioelectronic systems214 (Fig. 9c). However, a critical challenge for hydrogel-based electronics is their susceptibility to dehydration in ambient air, which can compromise their long-term stability and performance. A feasible solution is to introduce water retention agents. For instance, with the addition of glycerol, the gelatin-based hydrogel electrodes remain stable for several days215, even with an ultrathin geometry112 (Fig. 9d).

Fig. 9: Soft electronics based on hydrogels.
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a Ultrathin hydrogel substrate for skin-conformable electronics. Reproduced with permission from ref. 212. Copyright 2023 Wiley-VCH. b PEDOT:PSS based conductive hydrogel with on-demand adhesion. Reproduced with permission from ref. 213. Copyright 2023 Springer Nature. c Self-healable hydrogel electrode based on polyvinyl alcohol-borate hydrogel and carbon nanotubes. Reproduced with permission from ref. 214. Copyright 2017 Wiley-VCH. d Nanomesh hydrogel skin sensor for long-term monitoring. Reproduced with permission from ref. 112. Copyright 2024 American Association for the Advancement of Science.

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Liquid metals

Recent advances have highlighted the exceptional potential of liquid metals for soft electronics133,134,216,217,218,219,220, owing to their unique combination of metallic conductivity and fluidic behaviour. Among the various liquid metals, eutectic gallium indium alloy (EGaIn, \({T}_{\text{melt}}\text{=}15.5^\circ \text{C}\)) has been extensively studied for its simple composition, excellent biocompatibility, low vapour pressure, commercial availability, and room-temperature liquidity134, which enables innovative applications in stretchable bioelectronics. For example, Zhuang et al. adopted EGaIn as a stretchable wiring and solder for a 3D integrated electronic skin95 (Fig. 10a), which enables good contact of electronic components under extreme strains of up to 1500%. In addition, by applying ultrasound and surfactants in water, EGaIn can be dispersed in water to form a printable ink, which supports the fabrication of flexible electronics by various techniques such as spray coating13 (Fig. 10b), direct writing on the skin surface221 (Fig. 10c), and 3D printing222,223. Notably, by controlling the oxidation degree of the liquid metals, the surface tension can be tuned, offering handles to control the fluidity, morphology, and processability224. For example, introducing partial Ga oxidation improves the adhesion of liquid metal225, enabling direct painting of liquid metal based skin electrodes (Fig. 10d).

Fig. 10: Soft electronics based on liquid metals.
figure 10

a Stretchable integrated circuits based on liquid metal wiring and solder. Reproduced with permission from ref. 95. Copyright 2024 Springer Nature. b Liquid metal ink for spray-coating fabrication of skin electronics. Reproduced with permission from ref. 13. Copyright 2024 Elsevier. c Liquid metal ink for direct painted skin sensor. Reproduced with permission from ref. 221. Copyright 2022 Wiley-VCH. d Skin electronics painted on skin with partially oxidised EGaIn alloy. Reproduced with permission from ref. 225. Copyright 2019 Wiley-VCH. e Breathable elastic conductor based on nanofibre mat coated with liquid metal. Reproduced with permission from ref. 111. Copyright 2021 Springer Nature.

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While liquid metal-based conformable bioelectronics offer remarkable functionality, one of their significant challenges is the non-permeability226, which may deteriorate normal physiological processes such as sweating. To solve this problem, researchers deposited liquid metal on a non-woven fibre mat111 (Fig. 10e), which was stretched to form a permeable porous structure before usage. The encapsulation of liquid metal should also be considered in some scenarios to prevent leakage227,228. Additionally, because of the low abundance and critical roles in other semiconductor applications of Ga, the recyclability of liquid metal-based bioelectronics should be considered to reduce costs and improve sustainability207,229.

Interfacial adhesion enhancement

Strong interfacial adhesion plays a critical role in achieving and maintaining conformal contact between bioelectronic devices and biological tissues28,230,231, because adhesion provides a negative energy contribution to the conformal process and ensures device stability during dynamic physiological movements. Although van der Waals (vdW) forces are universally present between contact surfaces, they are generally insufficient to ensure robust adhesion for bioelectronics232. Conventional methods, such as suturing, medical glue, and medical tape28, have been leveraged for stable attachment. However, conventional methods usually compromise conformability and/or damage biosurfaces; therefore they are not suitable for next-generation bioelectronics. To address the limitations of conventional fixation methods, various interfacial adhesion enhancement strategies have been developed.

Microstructured surface

Recent advances in bioelectronics have incorporated various microstructural designs to enhance interfacial adhesion at the bioelectronic–tissue interface. Among these, bioinspired approaches have proven to be particularly effective233,234,235,236,237,238. The gecko-foot-inspired micropillar array is one of the most successful designs233,239,240,241, where the discrete contact points of the array enable conformal adaptation to rough biosurfaces while significantly increasing the total adhesion force. Kim et al. designed a self-adhesive dry electrode with a gecko foot-inspired microstructure for stable electrocardiogram (ECG) recordings183 (Fig. 11a). With a micropillar array replicated from the silicon wafer mould, such electrodes achieve a high adhesion force of 1.3 N/cm2 with skin, which enables conformable mounting of the electrode on the skin and stable ECG recordings under various movement scenarios. Another promising adhesive microstructure is the octopus-inspired microsucker array design242,243,244,245, which generates an adhesion force based on the pressure difference between the inside and outside of the sucker. Chun et al. introduced an octopus-inspired microsucker structure into an elastic conductive film to fabricate skin-attachable biosensors242 (Fig. 11b). Such skin conformable device realises a high adhesion with skin (1.89 N/cm2), enabling stable measurement of joint bending and ECG even underwater.

Fig. 11: Interfacial adhesion enhancement.
figure 11

a Dry adhesive skin electrodes based on micropillar array. Reproduced with permission from ref. 183. Copyright 2016 American Chemical Society. b Adhesive skin electronics based on microsucker structure. Reproduced with permission from ref. 242. Copyright 2018 Wiley-VCH. c Adhesion mechanism of mussel’s byssus. Reproduced with permission from ref. 253. Copyright 2021 Wiley-VCH. d Wireless optoelectronics with tissue adhesiveness via polydopamine coating. Reproduced with permission from ref. 258. Copyright 2019 Springer Nature. e Dry double-sided tape with strong tissue adhesion via chemical bonding and its application in implanted conformal sensor on heart surface. Reproduced with permission from ref. 249. Copyright 2019 Springer Nature. f Conductive bioelectronic interface with strong tissue adhesion via chemical bonding. Reproduced with permission from ref. 260. Copyright 2021 Springer Nature. g bioadhesive semiconductor material for implanted conformal bioelectronics. Reproduced with permission from ref. 261. Copyright 2023 American Association for the Advancement of Science.

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Adhesive material with interfacial bonding

Enhancing intrinsic interfacial interactions significantly improves adhesion230,231, which is closely related to the chemical structure of the material in contact with the biosurfaces. Typically, adhesion interactions between the biosurface and bioelectronics are dominated by non-covalent interactions such as vdW interactions, which are naturally weak. For example, the adhesion between common PDMS substrate and skin surface is ~0.2 N/m48,246. Although stronger non-covalent interactions have been introduced to enhance the bioelectronic–biosurface interfacial adhesion164,247,248, such improvements are usually limited because the typical bonding energy of non-covalent interactions is lower than 10 kcal/mol. Adhesive materials with chemical bonds have been incorporated in bioelectronics to further improve the interfacial adhesion230,249,250,251,252.

Nature provides an exemplary model of strong interfacial bonding in the marine mussel’s byssus, which achieves remarkable adhesion through catechol-rich proteins253,254 (Fig. 11c). Detailed studies have revealed that catechol groups form diverse interactions with surface sites255,256,257, including multiple hydrogen bonds, coordination bonds with metal ions, and covalent bonds with amine groups. The in-situ polymerisation of dopamine is an effective strategy for introducing mussel-mimetic adhesives because the as-produced polydopamine (PDA) thin layer contains numerous catechol groups254. For example, the PDA modification of the PDMS membrane improves its adhesion to muscles (Fig. 11d), showing a maximum separation force five times greater than that of the pristine PDMS membrane258. Such adhesive PDMS membranes ensure the stable conformal attachment of implanted wireless-powered light-emitting diodes to tissues, enabling the metronomic photodynamic therapy of tumours in animal models.

Rationally designed active reaction sites have been implemented on the surfaces of bioelectronics to improve the stability and controllability of interfacial adhesion. For example, the N-hydroxysuccinimide (NHS) group reacts with the amine group on the tissue surface, resulting in a stable amide group that enhance the interfacial bonding strength230,250,252,259. Yuk et al. developed a dry double-sided tape based on poly(acrylic acid) modified with NHS groups249 (Fig. 11e), which shows high reactivity with abundant primary amine groups on the surface of biological tissues. The interfacial covalent bonding provides a high fracture toughness of more than 1000 J/m2, which enables the stable attachment of strain sensors to the surface of the heart to monitor heartbeats. Moreover, the introduction of conductive graphene to such system achieves electric bioadhesive with a combination of good conductivity of 2.6 S/m and high interfacial toughness of ~400 J/m2 260 (Fig. 11f). Such an adhesive and conductive interface material enable the stable transmission of electrical signals between the biosurface and bioelectronics, facilitating biosignal recording and tissue stimulation. By combining the NHS-functionalized brush polymer and semiconducting polymer, Li et al. realised bioadhesive soft semiconducting film261 (Fig. 11g), which enables the fabrication of an OECT for the in-vivo amplification of electrophysiological signals.

Gaussian curvature mismatch reduction

The Gaussian curvature mismatch between bioelectronic and non-developable biosurfaces leads to a fundamental challenge in achieving conformal interfaces, as discussed in the section on global conformability on non-developable surfaces. Various structural design and material engineering strategies have been adopted to reduce the elastic energy accumulation and improve the bioelectronic–biosurface interfacial interaction, as shown in the sections on softness improvement and interfacial adhesion enhancement. However, such strategies fail to avoid the in-plane deformation introduced by geometry mismatch, which may compromise both device performance and tissue integrity during long-term use. Various strategies have been developed to reduce geometric mismatches.

Subdivision

As discussed in the section on global conformability on non-developable surfaces, the conformability of bioelectronics exhibits an inverse relationship with the device lateral dimensions54, making size miniaturisation a feasible choice for improving integration with non-developable biosurfaces. For instance, a critical size exists when a film is mounted on a spherical surface53,59. Additionally, reducing the lateral size decreases the in-plane strain54. To balance the conformability and spatial coverage of bioelectronics on large non-developable biosurfaces, subdivision strategy was developed. For example, to achieve the conformal mounting of a flexible photodetector array on a spherical substrate, a flattened truncated icosahedron geometry was adopted56, which is inspired by the soccer ball (Fig. 12a). The subdivision design successfully releases strain and avoids wrinkles, resulting in a human-eye-inspired hemispherical imager. Researchers have further explored various geometric adaptations such as triangle-mesh configurations262, edge-cutting patterns53, and petal-like (Fig. 12b)263,264 designs to reduce strain. Computational algorithms have been developed to generalise these subdivision approaches for arbitrary non-developable surfaces262 (Fig. 12c). While effective, these geometric solutions exhibit certain limitations. This approach requires precise a priori characterisation of the target substrate morphology to inform the device subdivision design. Furthermore, the necessary segmentation introduces discontinuities in the final device architecture, which may constrain the electronic functionality.

Fig. 12: Gaussian curvature mismatch reduction strategies.
figure 12

a Human eye inspired hemispherical imager achieved by truncated thin film optoelectronics. Reproduced with permission from ref. 56. Copyright 2017 Springer Nature. b 3D curved electronics by subdivision with petal-like design. Reproduced with permission from ref. 263. Copyright 2022 Wiley-VCH. c Computer assisted design of 3D surface subdivision. Reproduced with permission from ref. 262. Copyright 2020 American Association for the Advancement of Science. d Transfer printing of electronics on 3D curved surfaces by ballon stamp. Reproduced with permission from ref. 267. Copyright 2019 Springer Nature. e Transfer printing with liquid metal stamp. Reproduced with permission from ref. 271. Copyright 2024 Springer Nature. f Direct electrode fabrication on skin via adaptive 3D printing. Reproduced with permission from ref. 275. Copyright 2018 Wiley-VCH. g Direct writing of skin electrode. Reproduced with permission from ref. 272. Copyright 2022 Wiley-VCH. h Fibre electronics directly deployed on biosurfaces via orbital spinning. Reproduced with permission from ref. 276. Copyright 2024 Springer Nature.

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Adaptive fabrication

In contrast to conventional techniques that frequently fabricate bioelectronics on flat substrates, transfer printing enables the fabrication on curved substrates265,266. During a typical transfer printing process, microelectronic components are picked from the donor substrate using a stamp and printed on the receiver substrate by applying the stamp to the target surface. Notably, soft stamps are typically leveraged, allowing the transfer of electronics onto curved surface246,266,267,268,269. Since the microelectronics are placed directly on the biosurface without a mediating substrate, they remain decoupled from each other, reducing the mechanical stress from the geometric mismatch. For example, substrate-less bioelectronics can be transferred to the human body using a PDMS stamp270, resulting in highly conformable and flexible epidermal electronics. Sim et al. developed balloon-mediated transfer printing on 3D surfaces267 (Fig. 12d). As the balloon exhibits excellent deformability, it can deliver microelectronics to highly curved surfaces with a relatively low applied pressure. Furthermore, liquids can also mediate transfer printing, which offers the potential for complex surfaces for the fluidity of liquids. To address the challenge of low adhesion between the liquid medium and bioelectronics, Shi et al. employed melted Ga as a liquid stamp271 (Fig. 12e). The interaction between Ga and microdevices can be tuned by controlling the melting, enabling the effective pick-up by solidified Ga and damage-free deployment by melted Ga.

Conformable bioelectronics on complex biosurfaces have also been achieved via in-situ printing with functional inks272,273,274. As the functional ink follows the shape of the biosurface and fills the rough surface, the in-situ printing process ensures good conformability of the as-prepared bioelectronics on arbitrary surfaces. To achieve high precision patterning, Zhu et al. developed a close-loop feedback control system for the in-situ 3D printing on human skin275 (Fig. 12f). Ershad et al. developed a fully biocompatible ink composed of silver flakes and PEDOT:PSS273, which can be applied directly to the human skin to fabricate on-skin electronics (Fig. 12g). The in-situ fabricated electronics are conformally attached to the skin, which ensures the stable and high-quality recording of electrophysiological signals. Recently, Wang et al. developed a conductive gel based on gelatin, citrate salt, and glycerol215, which exhibits a reversible phase transition from a solid-like state to a liquid-like state during heating. The temperature-sensitive rheology of biogel enables painting at higher temperature and in-situ gelation on the skin to form skin electronics. Notably, direct painting enables conformal skin electrodes to be placed on hairy surfaces such as the scalp44,215, allowing high-quality electroencephalogram (EEG) measurements without skin preparation.

Fibre electronics

Fibre electronics hold significant potential as conformable bioelectronics on curved surfaces given their good flexibility. Moreover, the 1D geometry of fibre cancels out the possible developability mismatch between bioelectronics and biosurfaces, ensuring conformability on arbitrary surfaces. Similar to thin-film devices, the high flexibility and adhesion promote the conformability of fibre electronics on curved surfaces. For example, the spinning of a PEDOT:PSS dope results in ultrathin conductive fibre276,277. By leveraging an orbit-spinning configuration, such conductive fibre was directly deposited onto various biosurfaces such as fingertips, chicken embryos, and orchid flower petals (Fig. 12h). Moreover, the residual water improves the flexibility and surface adhesion of the conductive fibre, further improving its conformability. Diverse conformal bioelectronics, such as physiological sensors, touch sensors, and OECT, can be realised using this technique.

Conclusion and outlook

The seamless integration of bioelectronics and biosurfaces has significantly advanced interdisciplinary fields bridging biology and electronics, resulting in significant academic and commercial value. In this review, we summarise the typical strategies for achieving conformal bioelectronics supported by theoretical models and experimental validation. First, the incorporation of flexible geometries and soft electronic materials reduces the rigidity of bioelectronics, enabling bending and stretching deformations to adapt to curved biosurfaces. Second, enhanced bioelectronic–biosurface adhesion offers driving force for conformability, ensuring stable and robust interfaces. Additionally, geometric mismatch has to be addressed when deploying bioelectronics on non-developable biosurfaces, as an incompatible Gaussian curvature can result in stress and wrinkles that compromise device performance.

Among the three strategies for achieving conformability on biosurfaces, softness improvement reduces the stress induced by deformation, thereby minimising the tissue impact and user awareness. However, this approach often sacrifices the electrical performance, durability, and ease of handling. Adhesion enhancement secures stable bioelectronic–biosurface interfaces, enabling long-term stability under dynamic conditions, but requires carefully engineered adhesive structures or materials, with biocompatibility remaining a critical consideration. Gaussian curvature mismatch reduction effectively improves conformity to complex, non-developable biosurfaces, yet such designs are often incompatible with conventional electronic fabrication processes, potentially limiting adaptability and scalability.

These insights lay the foundation for future innovations in bioelectronics, fostering tighter integration with biological systems for advanced applications. Despite these advances, the inherent complexity of biosurfaces necessitates more sophisticated designs for advancing conformal bioelectronics.

Conformal bioelectronics on hairy surface

Hairy structures are ubiquitous across biosurfaces, including humans13,44,243, animals278,279, insects45,46, and plants22,280, where they perform critical functions such as thermal regulation, sensing, protection, and water manipulation. However, these structures pose significant challenges for conformal bioelectronics by impeding adhesion and creating interfacial voids, ultimately degrading the performance29,31. For instance, high-quality electrophysiological recording on hairy skin often requires shaving the target site or applying a conductive paste to ensure proper electrode contact30. To overcome this limitation, electrodes with pillar arrays have been developed for well contact between the electrode and skin through hair243,281. However, these electrodes are generally bulky and uncomfortable. Emerging in-situ fabrication techniques using liquid precursors offer a possible solution for this challenge. For example, printing of conductive inks enables conformal electrode deposition on hairy skin for stable and high-quality electrophysiological recordings44 (Fig. 13a). Temperature-induced gelation has also been leveraged to fabricate conformal electrodes on the human scalp215 (Fig. 13b) or hairy stem of plants22 (Fig. 13c) to record electrophysiological signals. Moreover, temperature-sensitive gel electrodes can be easily removed using warm water without damaging the underlying skin215. Nevertheless, the current advancements are largely limited to the surface electrodes. Future studies should explore novel bioelectronic designs that seamlessly integrate with hairy biosurfaces to broaden the scope of wearable and implantable bioelectronics.

Fig. 13: Perspectives for future conformal bioelectronics.
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a In-situ fabrication of elastic electrodes on hairy scalp for stable electrophysiological signal recording. Reproduced with permission from ref. 44. Copyright 2023 Wiley-VCH. b Temperature-sensitive biogel for electroencephalogram signal recoding on hairy scalp. Reproduced with permission from ref. 215. Copyright 2022 American Association for the Advancement of Science. c Temperature sensitive gel electrodes for recording plants electrophysiological signal on hairy stem. Reproduced with permission from ref. 22. Copyright 2021 Wiley-VCH. d Soft-rigid interface based Au nanoparticle embedded in styrene-ethylene-butylene-styrene (SEBS) elastomer matrix. Reproduced with permission from ref. 145. Copyright 2023 Springer Nature. e Wireless body network based on stretchable integrated chip-free passive sensor. Reproduced with permission from ref. 146. Copyright 2019 Springer Nature. f Wireless contact lens for eye pressure and glucose monitoring. Reproduced with permission from ref. 286. Copyright 2017 Springer Nature. g Nanomesh pressure sensor for interference-free finger-tip pressure monitoring. Reproduced with permission from ref. 119. Copyright 2020 American Association for the Advancement of Science. h Thermal-wet comfortable skin sensor based on PEDOT:PSS coated silk nanofibre. Reproduced with permission from ref. 170. Copyright 2021 American Chemical Society. i 3D liquid diode for permeable skin electronics. Reproduced with permission from ref. 282. Copyrights 2024 Springer Nature.

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Interface with rigid circuits

Bioelectronics are usually soft to achieve compatibility with soft tissue and reduce motion artefacts. However, rigid components remain essential for complex functionalities such as signal processing, data transmission, and power management95,282,283. The interface between the bioelectronics and rigid modules is thus important for the entire system. However, the mechanical mismatch between soft bioelectronics and rigid modulus leads to unstable connections and even interfacial failure, which significantly affects the practical applications284. Recent advances have addressed this challenge through innovative interface designs. For instance, a universal soft-rigid interface was developed by embedding conductive metal nanoparticles into a SEBS elastomer matrix145 (Fig. 13d), offering high conductivity, excellent stretchability, and strong interfacial adhesion to seamlessly integrate soft bioelectronics, rigid electronics, and encapsulation layers. Gradient modulus designs have also been employed to reduce the mechanical mismatch at the interface115.

Wireless systems provide an effective alternative to eliminate physical connections and address interface mismatch challenges in bioelectronics. Notably, resistor-inductor-capacitor (RLC) resonant circuit represents a popular design for wireless bioelectronics285. The resonance frequency \({f}_{0}=1/2\pi \sqrt{{LC}}\) and quality factor \(Q=\sqrt{\frac{L}{{R}^{2}C}}\) of a radio-frequency RLC resonant circuit can be wirelessly measured by a readout circuit, enabling the contactless derivation of equivalent resistance (R), inductance (L), and capacitance (C). For instance, body-conformable stretchable and chip-free RLC resonant sensors with strain-responsive resistance have been developed146, allowing wireless monitoring of breathing, pulse, and body movement using flexible readout circuits integrated on cloth (Fig. 13e). Smart contact lenses embedded with RLC resonant sensors have realised wireless eye pressure monitoring286 (Fig. 13f), glucose level monitoring286,287, and on-demand drug release288,289. Wireless bioelectronics based on the electromagnetic coupling between the body and metamaterial textiles290,291 or fibre electronics292 have also been reported.

Interference-free bioelectronics

Biosurfaces inherently perform various critical physiological functions. For example, the human skin is considered the largest organ and possesses important functions of protection, thermal regulation, and sensation to multiple stimuli. An ideal skin-interfacing system should maintain these native functions and minimise user awareness7,32,293. Recent advances in conformal bioelectronics have addressed this challenge using innovative materials and structural designs. For instance, reducing device stiffness enhances mechanical compliance with the skin, mitigating loss of pressure sensation2,65,294. The nanomesh architecture further improves softness, enabling negligible interference with tactile perception, even on highly sensitive sites such as fingertips119 (Fig. 13g). Notably, the ultra-low stiffness of tactile perception-free bioelectronics leads to significant strain induced by normal physiological activities, external mechanical inference, or unintentional deformation during deployment process, resulting in challenges to robustness. Ultra-soft bioelectronics with balanced stiffness and robustness have been achieved by improving the interactions of conductive layer with ultrathin substrates118,295 or biosurfaces248,258.

Another key consideration is permeability, as impermeable devices disrupt the sweating process of the skin, which is vital for natural thermoregulatory and excretory functions296,297. Conventional non-breathable devices trap sweat at the skin interface, leading to interfacial delamination, impaired heat dissipation, user discomfort, and potential inflammation. To overcome this challenge, permeable structures such as ultrathin films, 2D or 3D porous structures, textile structures, and mesh structures have been introduced in conformal bioelectronics to improve permeability170,296 (Fig. 13h). However, current research on porous bioelectronics has primarily focused on conformal electrodes, leaving significant room for advancement in the development of complex integrated bioelectronic systems. Recently, liquid diode structure has been introduced to improve the transmission rate of sweat across skin electronics282, which provides a platform for multifunctional integrated bioelectronics (Fig. 13i). Moreover, optical influence-free bioelectronics have been fabricated based on transparent components21,176,293. Nevertheless, as the current study focuses only on limited aspects, achieving truly imperceptible bioelectronics in the future necessitates a systematic and integrated consideration of minimising interference across all aspects of skin physiology.

The function and conformability of bioelectronics can be further improved by incorporating novel materials, structures, and fabrication technologies. Future studies should consider other essential aspects to promote advancement, such as biological imperceptibility, system integration, fabrication efficiency, and sustainability. Moving forward, advancements in conformal bioelectronics are expected to drive major breakthroughs in fundamental science and practical applications.