Tissue-bioelectronics interfaces.

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Published in final edited form as: Chem Soc Rev. 2026 Jun 8;55(11):6307–6343. doi: 10.1039/d5cs01422k Tissue-bioelectronics interfaces Jiahong Li 1, † , Jin Qu 1, † , Wei Gao 1, * Author information Article notes Copyright and License information 1 Andrew and Peggy Cherng Department of Medical Engineering, Division of Engineering and Applied Science, California Institute of Technology, Pasadena, CA, USA † These authors contribute equally to this work. * Email: weigao@caltech.edu Collection date 2026 Jun 8. PMC Copyright notice PMCID: PMC13231841  NIHMSID: NIHMS2175218  PMID: 42108816 The publisher's version of this article is available at Chem Soc Rev Abstract Wearable and implantable bioelectronics enable continuous physiological monitoring and therapeutic modulation, yet their performance critically depends on the stability and conformality of the interfaces with soft and dynamic biological tissues. Mechanical and biochemical mismatches between conventional electronic materials and living tissues often lead to interfacial stress, unstable contact, and inflammatory responses that compromise the long-term function of bioelectronics. This review presents a mechanism-driven framework for understanding tissue–bioelectronics interfaces by systematically examining the physical, chemical, and biological interactions that govern device–tissue coupling across temporal and length scales. We show how these interfacial mechanisms inform key design principles in structural engineering and materials development, enabling improved mechanical compliance, adhesion, and long-term interfacial stability. We further highlight recent advances in fabrication strategies that support soft, conformal, and multifunctional bioelectronic systems, together with representative applications spanning physiological sensing and therapeutic modulation. Finally, we discuss emerging strategies for mitigating foreign-body responses and outline remaining challenges and opportunities for achieving durable, adaptive, and clinically translatable tissue–bioelectronics interfaces. Graphical Abstract A mechanism-driven framework for tissue–bioelectronics interfaces, integrating interfacial interactions, materials and structural design, and fabrication strategies to enable stable, conformal bioelectronic systems for continuous monitoring and therapy. 1. Introduction Driven by rising healthcare costs, the increasing prevalence of chronic diseases, and aging populations requiring long-term care, healthcare systems are increasingly incorporating distributed, continuous monitoring approaches enabled by wearable and implantable bioelectronic technologies, complementing conventional centralized and reactive care models 1 , 2 . Advances in wearable and implantable bioelectronics enable both continuous monitoring that allow early detection of health abnormalities and sustained therapeutic modulation 3 – 8 . The effectiveness of these technologies depends on the tissue–bioelectronics interface; thus, establishing a stable and conformal interface is vital for accurate biosignal acquisition and consistent therapeutic performance. A major challenge arises from the pronounced mechanical mismatch between conventional rigid bioelectronic devices and soft, irregularly shaped biological tissues. This mismatch induces interfacial stresses and motion-induced damage, which, together with biochemical incompatibility, can lead to tissue inflammation, fibrotic encapsulation, and premature device failure 9 – 12 . At the molecular level, interfacial stress and micromotion can activate cellular mechanotransduction pathways, including integrin–focal adhesion signaling and cytoskeletal remodeling 13 , 14 . These processes trigger immune activation and cytokine release (e.g., TNF-α, IL-1β, and IL-6), which recruit macrophages and fibroblasts to the interface 15 . Sustained inflammatory signaling and TGF-β–mediated pathways further promote extracellular matrix deposition and fibrotic capsule formation, progressively isolating the device from surrounding tissues and degrading signal coupling 9 . In addition to mechanical incompatibility, biochemical mismatch between synthetic materials and the native extracellular matrix can disrupt the interfacial microenvironment. Protein adsorption and biofouling on material surfaces often initiate immune recognition, leading to macrophage activation, cytokine release, and foreign-body responses that contribute to fibrotic encapsulation and long-term device failure 16 . At the cellular level, these perturbations can alter cell morphology, cytoskeletal organization, and adhesion signaling, influencing the behavior of immune cells such as macrophages and fibroblasts at the interface 17 . Persistent inflammatory signaling further promotes extracellular matrix (ECM) remodeling, including collagen deposition and fibrotic capsule formation around implanted devices 10 . These processes progressively modify the local microenvironment, increase interfacial impedance, and ultimately contribute to the foreign body response that isolates the device from surrounding tissues. Importantly, mechanical compliance at the device level alone does not guarantee functional stability, as both biosignal acquisition and therapeutic stimulation essentially rely on efficient coupling across the tissue–device interface 18 – 20 . To address these challenges, extensive efforts have been devoted to optimizing both device architectures and material compositions. Structural design strategies provide mechanical compliance by allowing rigid components to deform with the surrounding tissue 21 , 22 , while the soft and biologically compatible materials further enhance mechanical compatibility and interfacial stability 23 – 25 . Notably, device architecture and material composition are often intrinsically coupled, jointly governing interfacial stress distribution and contact mechanics. These approaches have substantially improved the robustness and longevity of tissue–device integration. However, interfacial stability is not a static outcome of material softness or structural compliance; rather it is a dynamic process shaped by mechanical loading, biochemical exposure, and biological remodeling over time. Achieving robust tissue integration therefore requires a comprehensive understanding of the interfacial interactions that govern contact, adhesion, and long-term stability at the tissue–device boundary. These interactions can be broadly categorized into physical, chemical, and biological mechanisms. Physical forces govern initial contact and conformability, chemical interaction—particularly covalent bonding—enhance interfacial toughness, and biological responses ultimately determine chronic biocompatibility. Understanding how these mechanisms operate across length and time scales is essential for designing interfaces that preserve both functional performance and tissue health. In this review, we summarize the fundamental interfacial mechanisms that govern tissue–device coupling. Adopting a mechanism-driven perspective, we then discuss key principles underlying structural designs and material innovations. We further review recent fabrication strategies enabling soft, conformal, and multifunctional bioelectronics, followed by representative applications in sensing, stimulation, and therapeutic systems. Finally, we outline emerging opportunities and remaining challenges that will shape the next generation of tissue–bioelectronics interfaces. Collectively, this review provides a framework that links interfacial mechanisms to material selection, structural design, and fabrication strategies for achieving stable tissue–device coupling. 2. Interfacial mechanisms At the device–tissue boundary, bioelectronic performance is governed by a set of interfacial mechanisms that determine how interfaces form, stabilize, and evolve over time. These mechanisms determine how effectively a device conforms to tissue, maintains adhesion, and preserves stable signal coupling over time under physiological conditions. Here we categorize interfacial interactions into three principal classes: physical interactions (e.g., van der Waals forces and capillary effects) 20 , 26 , chemical interactions dominated by covalent bonding between device surfaces and tissue components 27 , and biological interactions mediated by cell adhesion and immune signaling 28 . An overview of these interaction types is provided in Fig. 1 . By outlining these foundational processes, we establish a conceptual framework for the rational design of stable and long-lasting device–tissue interfaces. Fig. 1. Open in a new tab Interaction on tissue-bioelectronics interface. (a) Physical interactions. (b) Covalent interactions. (c) Biological interactions. 2.1. Physical interactions Physical interactions govern the formation of initial contact and conformal coupling between bioelectronic devices and soft biological tissues, spanning length scales from molecular-level surface forces to macroscopic structural engagement ( Fig. 1a ). At the microscale, weak noncovalent interactions, including van der Waals forces and hydrogen bonding, contribute to intimate surface contact and adhesion 26 , 29 . At larger length scales, capillary forces arising from interfacial fluids, as well as geometric and mechanical interlocking enabled by device architecture or tissue microtopography, can also enhance interfacial adhesion and load transfer 20 , 30 . Collectively, these physical interactions play a central role in minimizing interfacial gaps, achieving conformability, and stabilizing tissue–device interface under physiological deformation. Van der Waals interactions Van der Waals interactions arise from dipole-dipole interactions between adjacent matter and are ubiquitous when materials are brought into close proximity. Although individually weak (~0.5–5 kJ mol −1 ) 31 , their collective contribution can yield appreciable interfacial adhesion when conformal contact maximizes effective contact area, particularly in ultrathin and mechanically compliant devices. Skin-like electronic systems have demonstrated robust adhesion relying solely on van der Waals interactions, enabled by intimate conformal contact that maximizes the effective interfacial area and by their ultralight weight. Such epidermal electronic systems support high-fidelity recording of electrophysiological signals such as electrocardiogram (ECG), electroencephalogram (EEG), and electromyogram (EMG), as well as multimodal sensing of temperature, strain, and optical signals 26 . Importantly, van der Waals–mediated interfaces can support not only passive sensing but also the reliable operation of active electronic components such as transistors on soft and dynamic tissues. For example, stretchable electronic membranes incorporating molybdenum disulfide (MoS 2 ) nanosheet–based exhibit enhanced resistance to motion-induced artifacts compared with conventional gel-based Ag/AgCl electrodes, yielding improved signal-to-noise ratios under dynamic skin deformation 32 . Hydrogen bonds Hydrogen bonding arises from directional interactions between hydrogen donors and electronegative acceptors in neighboring molecules and is particularly prevalent at tissue–device interfaces due to the abundance of hydrogen-bonding functional groups in proteins. Weak hydrogen bonds (≈10–40 kJ mol −1 ) undergo rapid exchange and contribute limited interfacial stability in aqueous environments, where water molecules readily compete for bonding sites 33 . In contrast, stronger and more directional hydrogen bonds, such as those formed between carboxyl groups and amino acid residues (≈71–85 kJ mol −1 ) 33 , can provide a basis for more stable interfacial adhesion. When present at high density, hydrogen bonds enable efficient load sharing and continuous bond reformation at the tissue interface, collectively yielding strong and fatigue-resistant adhesion under wet and dynamic conditions 34 . Consequently, hydrogen-bond–based tissue adhesion strategies are widely employed in hydrogel-based tissue adhesives 29 , 34 – 37 . Electrostatic interactions Electrostatic interactions arise from Coulombic interaction between oppositely charged species and are common at tissue–device interfaces due to the abundance of charged functional groups in biological tissues 38 . Several studies have demonstrated strong tissue adhesion mediated by electrostatic interactions 39 , 40 . However, electrostatic interactions are often screened by dissolved ions in physiological environments, which substantially limits their contributions in long-term adhesion when used in isolation. Chain entanglement Chain entanglement provides a noncovalent yet mechanically robust mechanism for tissue adhesion through interpenetration and topological constraints of long polymer chains at an interface. The process typically involves diffusion of polymer chains or monomers into the tissue surface, followed by network formation and solidification that locks the matrix in place. Chain entanglement-based adhesion is particularly effective at overcoming interfacial hydration layers and accommodating irregular and rough tissue surfaces, and is therefore widely adopted in many glue-like tissue adhesive systems 41 – 43 . Capillary adhesion Capillary adhesion, governed by liquid-mediated capillary forces that develop at two solid interfaces through meniscus formation in the presence of a thin fluid layer, is frequently exploited to stabilize and fix bioelectronic devices, as internal organs are intrinsically hydrated. Under thin-film approximation, the capillary pressure scales inversely with the interfacial separation and depends on the combined wettability of the two surfaces 44 , 45 : Δ P = γ c o s θ 1 + c o s θ 2 h (1) where γ is the surface tension of the liquid, θ 1 and θ 2 are the contact angles on the two surfaces, and h is the interfacial gap thickness. Because the surface tension of biological fluids and tissue wettability are largely fixed under physiological conditions, engineering strategies that reduce the device contact angle and minimize the thickness of the interfacial hydration layer are effective routes to enhance capillary adhesion. The resulting adhesion energy generated by capillary forces typically falls in the range of ~10–100 mJ m −2 , making this mechanism particularly well suited for thin, flexible implantable electronic systems 20 , 46 – 48 . Mechanical fixation Mechanical fixation stabilizes tissue–device interfaces through geometric interlocking, friction, or physical confinement, and is inherently insensitive to surface chemistry, hydration layers, or ionic environments. Representative strategies include sutures, staples, barbs and microneedles provide robust fixation under large deformation and dynamic loading, but often induce tissue damage. Less invasive approaches, including suction-based attachment and mechanical interlocking structures, have been explored to stabilize devices under wet conditions 49 – 51 . However, their broader application to bioelectronics remains limited, due to the complex mechanical and biological characteristics of living tissues. 2.2. Covalent interactions Covalent interactions establish strong linkages between bioelectronic devices and biological tissues, with bond dissociation energies typically in the range of ~100–800 kJ mol −127 , substantially exceeding those of noncovalent physical interactions ( Fig. 1b ). Owing to their short interaction range, covalent bonds require intimate molecular contact and precise chemical complementarity between functional groups on device surfaces and reactive moieties within tissue constituents. In physiological environments, blood and interstitial fluids present on the tissue surface may foul device surfaces or compete with tissue functional groups, thereby compromising effective covalent bonding 27 . To overcome these challenges, covalent interactions are typically combined with strategies that disrupt interfacial hydration layers and with physical interactions that promote rapid conformal coupling and molecular proximity between devices and soft tissues 35 , 39 , 42 . Within this framework, a range of covalent chemistries has been developed to enable tissue–device bonding by targeting functional groups common to both bioelectronic materials and biological tissues. N-hydroxysuccinimide esters N-hydroxysuccinimide (NHS) esters are among the most widely used chemistries for protein modification and bioconjugation 52 . This reactivity has been adapted for hydrogel-based adhesive systems, in which NHS ester groups incorporated into polymer networks form covalent bonds with tissue amine- and thiol-containing groups on tissue under physiological conditions 35 , 43 , 53 . In addition to mediating tissue bonding, NHS esters are frequently employed as crosslinkers to generate covalently bonded polymer networks, producing bulk adhesive matrices that form the basis of several commercial medical glues 54 , 55 . Despite their versatility, NHS esters are susceptible to hydrolysis, which poses challenges for long-term storage and reduces bonding efficiency in persistently wet conditions. Aldehydes Aldehyde-based chemistries have also been extensively explored as reactive platforms for covalent tissue adhesion owing to their ability to form Schiff base linkages with amine groups on biological tissues and to induce covalent crosslinking within polymer matrices 56 , 57 . This dual functionality underlies the design of commercially available medical glues in which glutaraldehyde-induced crosslinking generates a robust bulk adhesive matrix while anchoring to tissue surfaces 58 . However, Schiff base bonds are dynamic and relatively low in bond stability 59 , and residual aldehyde species can raise concerns regarding cytotoxicity and long-term biocompatibility. Cyanoacrylates Cyanoacrylates represent one of the most widely used classes of chemical adhesives, owing to their rapid bonding kinetics and strong adhesion. In fact, household “super glue” formulations are based on cyanoacrylate chemistry. Cyanoacrylate adhesion relies on anionic polymerization that is initiated by trace amounts of moisture, amines or other nucleophilic species present on surfaces, leading to rapid in situ polymerization and the formation of solid polymer networks that anchor to the substrate at the interface 60 , 61 . Short-chain cyanoacrylates, such as methyl and ethyl cyanoacrylate, exhibit extremely rapid polymerization kinetics and release substantial heat during curing. In wet biological environments, these monomers often react prematurely with interfacial water, limiting effective bonding to underlying tissue surfaces. To overcome these challenges, commercial medical glues employ cyanoacrylate monomers with longer alkyl side chains to deliberately moderate polymerization kinetics and improve handling and biocompatibility in physiological environments 62 – 64 . Nonetheless, the relatively fast degradation of cyanoacrylate-based adhesives generally restricts their use to short-term applications 65 . Catechol Catechol-based chemistries, inspired by mussel adhesion, have been widely exploited in biomimetic adhesive systems 66 – 68 . Catecholic moieties support multimodal adhesion by enabling physical interactions, such as hydrogen bonding and π–π interactions, as well as multiple covalent reaction pathways with nucleophilic groups on tissue components 69 . When oxidized to quinones, carbonyl groups can form Schiff base linkage with amines, while benzene ring can form covalent interaction with amines, thiols and imidazole 70 . While catechol offers versatile and strong tissue adhesion, its susceptibility to uncontrolled oxidation and self-crosslinking complicates regulation of quinone reactivity, posing challenges for practical and reproducible long-term applications. Other functional groups Other reactive functional groups, such as isocyanates, and aryl azides, have also been explored for covalent tissue adhesion, owing to their strong electrophilicity and ability to react with nucleophilic groups on biological tissues 71 , 72 . However, these chemistries remain relatively underutilized in tissue–bioelectronics applications, owing to concerns related to cytotoxicity, nonspecific reactivity, limited stability in wet physiological environments, and practical challenges such as the need for additional activation steps or specialized reaction conditions, which complicate their use in clinical settings. 2.3. Biological interactions Biological interactions govern the long-term fate of bioelectronic devices by mediating dynamic and reciprocal processes between implanted materials and living tissues at the tissue–device interface ( Fig. 1c ). Unlike physical and chemical interactions that primarily determine initial contact and adhesion, biological interactions operate over extended time scales and shape chronic biocompatibility by regulating immune recognition, cellular behavior, and tissue remodeling at the cellular and molecular levels. While foreign body reactions—characterized by macrophage activation, foreign body giant cell formation, and fibrotic encapsulation—represent a prominent and often adverse biological response to implanted devices 15 , biological interactions extend beyond these defensive processes to include adaptive and integrative responses such as cell adhesion, migration, differentiation, and functional tissue integration 28 . Collectively, biological interactions determine whether bioelectronic interfaces are merely tolerated, actively isolated, or ultimately integrated with host tissues. Passive interfaces Passive interfaces aim to improve long-term tissue compatibility by minimizing biological perturbation at the tissue–device interface. These strategies rely primarily on biomimetic mechanical and structural design principles, including soft materials, miniaturized dimensions, and tissue-matched geometries, to reduce mechanical mismatch and foreign-body recognition following implantation. By limiting the magnitude of injury, micromotion-induced stress, and abnormal force transfer, passive interfaces can attenuate immune activation and delay the onset of chronic inflammatory responses 73 , 74 . Ultra-thin polymer–metal–polymer architectures exemplify this approach, as their extremely low bending stiffness enables close mechanical matching with soft tissues and has yielded encouraging performance in long-term implantation studies 75 – 77 . More recently, hydrogel-based electronics have attracted increasing attention, as their intrinsically soft, hydrated, and viscoelastic nature closely resembles that of biological tissues, further promoting mechanical compatibility and interfacial stability 78 – 80 . Bioactive interfaces Bioactive interfaces refer to engineered tissue–device interfaces that actively interact with and regulate the biological microenvironment through biochemical functionalities. By incorporating bioactive cues, such as extracellular matrix proteins, adhesion molecules, and growth factors, these interfaces are designed to modulate host responses such as immune activation, inflammation, and tissue remodeling, rather than relying solely on passive mechanical accommodation 81 – 83 . In addition to immunomodulation, bioactive interfaces can enhance selective adhesion of target cells by presenting specific biochemical signals that promote controlled cell–material interactions 82 , 84 . Through these mechanisms, bioactive strategies can influence macrophage polarization, suppress fibrotic capsule formation, and enhance cell adhesion at the interface, thereby improving longterm functional stability and signal fidelity of implanted bioelectronic devices. Biohybrid interfaces Biohybrid interfaces represent a more integrative strategy in which living biological components—such as cells or tissues—directly with electronic systems to establish dynamic, bidirectional interactions 28 . By leveraging the inherent adaptability of living tissues, biohybrid interfaces offer unique opportunities for achieving long-term functional integration and adaptive performance. For example, the use of a sciatic nerve fragment as a biohybrid electrode has enabled stable long-term neural recordings with minimal foreign body response following chronic implantation 85 , 86 . Similarly, cell-seeded electrodes have shown improved performance in neural recording and neural regeneration by using living cellular components to promote tissue integration and functional coupling at the interface 87 – 89 . However, a key limitation of biohybrid approaches is the extended time—often weeks to months—required stable physiological interactions with host tissues, which remains a barrier to rapid functional deployment 28 . 3. Principles for conformal interface on irregular surfaces While the interfacial mechanisms discussed above elucidate how tissue–device interactions arise at the molecular, cellular, and tissue levels, effective bioelectronic integration ultimately requires precise control over physical contact between devices and complex biological surfaces. The stability of such conformal interfaces is governed by the interfacial adhesion strength and the mechanical rigidity of the device. Whereas strategies to enhance adhesion strength were addressed in Section 2 , this section focuses on fundamental design principles for reducing device rigidity to enable conformal contact. Here, we highlight two fundamental design strategies: structural engineering, which enables devices to geometrically adapt to tissue curvature and motion 22 , and materials development, which utilizes intrinsic material properties to match tissue softness, stretchability, and physiological environments 23 , 24 . Together, these principles establish a unifying framework for designing bioelectronic interfaces capable of maintaining robust and durable contact with irregular biological surfaces in both wearable and implantable applications. 3.1. Principles of structure engineering for conformal interfaces Structural engineering in bioelectronic interfaces refers to the deliberate design of device geometry, architecture, and layout to tailor effective mechanical behavior beyond intrinsic material properties. Among the relevant mechanical parameters, bending stiffness is a primary determinant of structural flexibility. For a rectangular thin-film or beam with width b and thickness h , the bending stiffness ( D ) is given by 90 : D = E b h 3 12 (2) where E is the Yong’s modulus of the material. The cubic dependence on thickness highlights the dominant role of thickness reduction in enhancing structural flexibility. Consequently, many structural strategies for flexible bioelectronics promote bending-dominated deformation. A representative example is the use of serpentine geometries, which transform axial deformation into bending and rotation of curved segments, thereby substantially reducing effective stiffness and material strain 91 . Under bending-dominated conditions, curved-beam mechanics suggests that the effective stiffness ( k e f f ) scales as 92 – 94 : k e f f ∼ E I R 2 L ∼ E h w 3 θ R 3 (3) where I is the second moment of area, R is the radius of curvature, and L is the arc length, h is beam thickness, w is width, and θ is the central angle. This scaling highlights the importance of large radius-to-width ratios in enabling high conformability and stretchability without excessive strain. Since biological tissues exhibit highly irregular, multiscale surface geometries, achieving conformal interfaces requires devices to be sufficiently flexible such that interfacial adhesion can overcome structural stiffness. As a simplest model, consider a rectangular thin film with Young’s modulus , thickness , and total length , conformally wrapped around a cylindrical substrate of radius ( Fig. 2a ), the required adhesion energy in unit area ( γ ) can be expressed as 20 : γ ≥ E h 3 24 R 2 (4) Fig. 2. Open in a new tab Representative structural engineering models and strategies for conformal tissue–bioelectronics interfaces. (a-d) Representative conformal bioelectronic–tissue interface geometries. a, Rectangular thin film conforming to a cylindrical substrate. b, Rectangular thin film adhered to partially overlapped cylinders. c, Circular thin film conforming to a hemispherical substrate. d, Circular strip conforming to a hemispherical substrate. (e) Representative structural engineering strategies. For a representative material (E ≈ 3 GPa) conforming to a cylinder (R ≈ 1 cm) via capillary interaction (γ ≈ 10 mJ m −2 ), the film thickness must be below ~20 μm. To model more complex geometries such as brain gyri, a pair of partially overlapping cylinders connected by a smooth arc can be considered ( Fig. 2b ). In this case, conformal adhesion requires the film thickness h to satisfy 20 : R s i n θ d − R s i n θ + d R c o s θ ( d − R s i n θ ) 2 − 24 γ R 2 E h 3 1 − d R s i n θ + d θ c o s θ R s i n 2 θ = (5) with θ = s i n − 1 d R + r (6) Using human cortical parameters (R ≈ 6.14 mm, d ≈ 5.93 mm, r ≈ 1.72 mm), the required thickness decreases to ~4.9 μm under capillary adhesion 20 . While cylindrical surfaces are developable and can, in principle, be conformally wrapped by thin films without inducing in-plane strain, many organs—such as the brain, heart, and stomach—exhibit non-developable geometries. For a planar circular thin film of radius r placed on a hemispherical surface of radius R ( Fig. 2c ), the circumferential mismatch between the film and substrate inevitably induces out-of-plane deformations such as wrinkling or buckling. However, when the contact area is sufficiently small ( θ ≪ 1 ), this mismatch becomes negligible and conformal contact can be achieved. An alternative solution for non-developable surface is to replace continuous films with narrow strip geometries ( Fig. 2d ). In this case, the required adhesion energy can be expressed as 20 : γ = E I R 2 + w 2 E h 24 r 2 1 − 1 − r 2 R 2 2 (7) where w is the strip width. Thin and narrow strips dramatically reduce the adhesion energy required to achieve conformal contact with non-developable surfaces, making capillary adhesion sufficient even on complex surfaces 20 . Taken together, these analyses show that achieving conformal interfaces with biological tissues fundamentally requires devices to be both thin and spatially small. Thickness reduction minimizes bending stiffness through its cubic scaling, thereby lowering the energetic penalty associated with bending. Simultaneously, reducing the lateral dimensions limits geometric mismatch, allowing modest adhesion energies to overcome the residual mechanical resistance. Here, these geometric and structural characteristics are collectively described as the device form factor, which encompasses its thickness, lateral dimensions, feature size, and architectural layout, and governs its mechanical behavior and interfacial conformability. In this context, large form factor typically corresponds to devices with extended spatial coverage or continuous geometries, whereas small form factor refers to devices with reduced feature sizes or discretized architectures that enable enhanced conformability. Guided by these principles, a range of structural designs have been developed to enable conformal interfaces between bioelectronics and biological tissues ( Fig. 2e ). Ultrathin structures, with thicknesses on the order of tens of micrometers or less, substantially reduce bending stiffness and allow relatively adhesion provided by van der Waals interactions or capillary forces to be sufficient for achieving conformal interfaces with biological tissues 3 , 26 , 32 , 95 . Microstructure engineering enhances conformability by introducing local mechanical compliance, in particular, micro-scale features such as microwrinkles and microcracks enable thin films to accommodate large deformations and complex surface curvatures by redistributing strain and reducing effective stiffness without sacrificing material integrity 96 – 99 . Kirigami-inspired designs introduce programmable cuts into thin films that transform in-plane stretching and compression into out-of-plane rotation, bending, and buckling of discrete structural units, enabling bioelectronics to accommodate large deformations and complex, non-developable surface curvatures while maintaining mechanical integrity and functional continuity 100 – 102 . Island-bridge designs connect rigid functional components with highly compliant interconnects, preserving high electronic performance while enabling mechanical adaptability 103 – 105 . Mesh electrodes further minimizing bending stiffness and volumetric footprint, enabling intimate mechanical coupling with soft and dynamic biological tissues 20 , 82 , 106 – 108 . Although increasingly minimal form factors improve mechanical flexibility and interfacial conformability, they often trade off component density and fabrication complexity. 3.2. Principles of materials development for conformal interfaces Materials development provides a complementary route to conformal bioelectronic interfaces with soft and irregular biological tissues by enabling shape adaptation and stress accommodation through intrinsic deformation behavior. From a mechanics perspective, materials employed for conformal bioelectronic interfaces can be classified according to their constitutive deformation behavior into elastic, viscoelastic, and viscous regimes, each offering a distinct mode of conformal adaptation through reversible deformation, time-dependent stress relaxation, or flow. This classification highlights how different materials accommodate tissue deformation and maintain conformal interfaces under physiological conditions. Representative materials within these regimes include elastomers such as silicones and polyurethanes (elastic) 109 – 111 , hydrogels (viscoelastic) 37 , 43 , 78 , 112 , 113 , and paste-like or liquid materials such as conductive gels or liquid metals (viscous). Elastic materials ( Fig. 3a ), characterized by a dominant storage modulus ( G ′ ) compared to the loss modulus ( G ′ ′ ), store the majority of the mechanical energy induced by deformation in a recoverable elastic form and obey Hooke’s law, in which deformation ( ε ) is linearly and instantaneously proportional to the applied stress ( σ ) 114 : ε = σ E (8) where E is the elastic modulus. As a result, sustaining stress is required to maintain conformal contact. Reducing elastic modulus lowers the stress needed to accommodate tissue curvature, thereby mitigating interfacial stress and facilitating the formation of conformal interfaces at soft and irregular tissue surfaces. Fig. 3. Open in a new tab Material regimes governing conformal tissue–bioelectronics interfaces. (a) Elastic regime (G′ ≫ G″): Materials deform reversibly with limited energy dissipation, providing structural stability but generating higher interfacial stress when conforming to soft tissues. (b) Viscoelastic regime (G′ ≈ G″): Time-dependent stress relaxation enables gradual adaptation to tissue surface irregularities, reducing interfacial stress and improving conformal contact. (c) Viscous regime (G″ ≫ G′): Flow-dominated deformation allows near stress-free conformity to complex geometries, but lacks structural stability for long-term interfacing. (d) Representative material classes across these regimes, including elastomers, hydrogels, and paste-like or liquid materials, illustrating how material selection governs the balance between mechanical stability and conformability at the tissue interface. Viscoelastic materials ( Fig. 3b ), characterized by comparable storage and loss moduli ( G ′ ≈ G ′ ′ ), dissipate a significant fraction of the mechanical energy induced by deformation and exhibit time-dependent mechanical behavior. Notably, many soft biological tissues also display viscoelastic responses over physiologically relevant timescales 115 . Such behavior can be approximated by the Maxwell model 116 : d ε d t = 1 E d σ d t + σ η (9) where ε is the strain, σ is the applied stress, E is the elastic modulus representing elastic energy storage, η is the viscosity associated with viscous dissipation, and t denotes time. Consequently, time-dependent deformation allows viscoelastic materials to gradually adapt to the substrate geometry, leading to partial relaxation of interfacial stress and stabilization of a conformal interface. Viscous materials ( Fig. 3c ), characterized by a dominant loss modulus ( G ′ ′ ≫ G ′ ), dissipate mechanical energy primarily through viscous flow and exhibit negligible elastic energy storage. Their deformation is governed by a constitutive relation of the form 117 : σ = η d ε d t (10) where σ is the applied stress, ε is the strain, η is the viscosity, and t denotes time. Viscous materials readily flow to conform to complex substrate geometries, such that conformal interfaces can be maintained with negligible interfacial stress. However, due to the lack of elastic restoring forces, viscous materials have limited ability to maintain a stable shape, making them more suitable for short-term applications. Elastomers, owing to their low elastic moduli and large reversible deformability, are widely used in bioelectronics to form conformal interfaces with soft tissues. Among these materials, silicone rubbers are one of the most extensively adopted elastomers due to their low Young’s moduli (~10 kPa–1 MPa), which are comparable to those of many soft biological tissues, as well as their excellent chemical stability and established biocompatibility, enabling reliable long-term interfacial contact in implantable systems 96 , 109 , 110 , 118 . In addition to silicones, thermoplastic polyurethanes (TPU) and styrenic thermoplastic elastomers, such as poly(styrene–ethylene–butylene–styrene) (SEBS), poly(styrene–butadiene–styrene) (SBS), and poly(styrene–isoprene–styrene) (SIS), are also commonly employed at biointerfaces. Owing to their mechanical resilience and structural stability, elastomeric materials are widely used not only as substrates and encapsulation layers in wearable and implantable bioelectronics 19 , 26 , 119 , 120 , but also as stretchable matrices for conductive composites and compliant interconnects 121 , 122 . Their ability to sustain large reversible deformation while maintaining structural integrity makes them particularly suitable for devices that undergo repeated mechanical loading, where mechanical robustness and long-term durability are required 123 – 127 . Hydrogels represent a representative class of viscoelastic materials for conformal tissue interfacing, owing to their low effective moduli, high water content, and pronounced stress relaxation behavior. Unlike elastic elastomers, hydrogels undergo time-dependent deformation under sustained loading, allowing interfaces to gradually adapt to surface roughness and complex tissue geometries while partially releasing interfacial stress 128 . Their tissue-like mechanical properties and hydrated nature further promote intimate contact with biological surfaces. Moreover, hydrogels inherently support ionic conductivity due to their hydrated networks and can be further engineered to achieve electronic conductivity by incorporating conductive fillers or conductive polymers 37 , 78 , 80 , 112 . This enables the formation of conductive and adhesive interfaces for bioelectronic applications. In addition, the viscoelastic response of hydrogels can be tuned through polymer composition, crosslinking density, and network architecture, enabling control over relaxation timescales and interfacial mechanics. These features make hydrogels particularly attractive for applications requiring long-term, stable conformal contact, such as adhesive cardiac patch, bioresorbable adhesive and wearable ultrasound interface 37 , 43 , 78 , 112 , 113 . Paste-like and liquid materials occupy the fluid end of the interfacial material spectrum. Paste-like or gel-based materials, such as ultrasound gels and conductive paste, possess weak network structures that provide limited elastic resistance, allowing them to spread and conform to complex tissue geometries under applied pressure or shear, while retaining a degree of shape stability in the absence of sustained external forces 125 – 127 . Notably, such paste-like interfacial behavior is not limited to gel-based materials. Eutectic gallium–indium (EGaIn), although liquid at physiological temperatures, exhibits partial shape retention ability owing to the oxide layer, which imparts paste-like behavior and enables conformal bioelectronic interfaces for electrophysiological signals recording 123 , 124 . Moreover, these characteristics make paste-like and liquid materials particularly suitable for applications requiring rapid conformal contact and transient interfacing, such as acoustic coupling layers (e.g., ultrasound gels) for ultrasound-based technologies and electrolyte gels for dry electrode configurations in electrophysiological recording 67 , 102 . 4. Fabrication strategies for bioelectronics interfaces Manufacturing routes for tissue–bioelectronics interfaces are constrained by the limited thermal/chemical budget and dimensional instability of soft materials, and by the need to maintain reliable electrical function under wet and dynamic conditions. Importantly, fabrication choices do not only determine feature resolution and yield; they also shape interfacial mechanics (conformality and strain distribution), wet adhesion/permeability, biofouling susceptibility, encapsulation integrity, and long-term signal/actuation drift. Accordingly, fabrication and integration should be treated as an extension of interface design, as they directly determine conformality, wet adhesion, encapsulation integrity, and long-term stability at the tissue interface 129 , 130 . 4.1. A practical selection framework: from interface requirements to manufacturing choices Manufacturing routes can be selected along five practical axes: feature density and resolution (dense microelectrodes versus large-area patterns); compatibility with soft and wet materials (elastomers, hydrogels, bioactive layers); dimensionality and architecture (planar laminates versus 3D porosity/gradients); integration strategy (monolithic soft systems versus hybrid assembly with rigid modules); and translation constraints (repeatability, quality control, sterilization/storage, and scalable production). Wearables often prioritize breathability and motion robustness; implants prioritize micromotion mitigation, chronic drift control, and encapsulation reliability; ingestible/intraluminal devices are frequently dominated by sealing, exposure-window design, and biofluid-fouling tolerance. A comparative matrix ( Table 1 ) summarizes major manufacturing families and their tissue-integration trade-offs. Table 1. Representative device fabrication and integration techniques for achieving conformal tissue–bioelectronic interfaces. Techniques Representative methods Best-fit applications Interface advantage Advantages Limitations Planar thin-film microfabrication PVD/CVD (incl. conformal polymer barriers) + photolithography High-density electrodes; miniaturized sensors; precise multilayers Enables ultrathin laminates + strain-engineered layouts for stable coupling Highest resolution; mature QC Thermal/solvent constraints; cracking/delamination; encapsulation dominates drift Soft lithography / micromolding Mold casting; replica molding; microfluidic replication Elastomer/hydrogel microstructures; biofluid handling Direct patterning in soft/wet-compatible materials improves coupling Broad material compatibility Lower density; sealing/adhesion critical under wet cycling Transfer printing / indirect assembly Pickup/placement onto soft substrates; float-type assembly Hybrid high-performance films on compliant receivers Combines electronic quality with tissue-like mechanics High performance + compliance Alignment/defects; yield issues on hydrated receivers Laser micromachining Cutting/ablation/engraving Rapid geometry optimization; large-area structuring Fast iteration for contact/adhesion geometry Maskless; rapid Thermal damage/debris; reproducibility on soft/wet layers Inkjet / aerosol jet printing Digital deposition of conductors/functional inks Wearables; flexible electrodes; coatings on non-planar surfaces; textile patterning Low-temp processing compatible with soft/porous substrates; enables maskless patterning Scalable; maskless Ink rheology; clogging; wet adhesion and multilayer defects DIW / extrusion printing Shear-thinning inks; multi-material 3D deposition Hydrogel/ionotronic interfaces; 3D porous constructs; localized textile coatings 3D compliance/porosit y improves coupling and mass transport Multi-material; 3D Resolution limits; curing consistency; batch reproducibility Screen printing Stencil-based large-area patterning Low-cost large-area wearables; robust textile electrodes/interconnects Tolerant to rough/porous surfaces; robust manufacturing for epidermal interfaces Very scalable Coarse features; limited density; encapsulation needed for wet stability Photopolymerizatio n 3D printing DLP/SLA; 2PP microstructuring Complex 3D scaffolds; microstructured components; UV-curable local encapsulation Programmable mechanics/geometry Complex forms Residuals; resin stiffness; speed–volume tradeoff In situ / in vivo fabrication Light-/ultrasound-triggered; extrusion/droplet in situ Geometry-adaptive wet adhesion; site-specific constructs Curing at site enhances conformity and wet adhesion Minimizes handling In vivo QC; curing completeness; residuals; regulatory burden Hybrid integration & packaging Layer stacking; adhesive bonding; connectorization Multifunctional systems; closed-loop platforms Mechanical decoupling + robust barriers reduce drift Preserves chip performance Encapsulation/connector failure; moisture ingress; fatigue Thermal drawing (preform-to-fiber) Preform design/assembly → thermal consolidation → drawing (tension/temperature control) → post-draw opening/exposure of channels/electrodes → terminal connectorization/strain relief Deep-tissue and intraluminal navigability; distributed multimodal fibers (electrical/optical/fluidic) Deterministic cross-sectional architecture + low bending stiffness enable minimally invasive access while maintaining function distribution along length Continuous long-length manufacturing; embedded multifunctionality; repeatable internal geometry Material co-draw constraints; post-draw access/exposure steps; terminal connectorization and wet packaging dominate chronic reliability Wet spinning & melt spinning (extrusion-based spinning) Spinneret extrusion of dope/melt → coagulation (wet) or thermal solidification (melt) → mechanical drawing/orientation → washing/annealing/crosslinking → optional coating/metallization + localized encapsulation Soft conductive/ionic fibers; biodegradable/suturable leads; wearable yarn electrodes; compliant routing where permeability/compliance matters High materials freedom enables tissue-mimetic mechanics and permeability; scalable fiber lengths for conformal coupling and routing Scalable and low-cost; broad material compatibility; tunable mechanics (draw ratio/orientation) Wet-state electrical stability and drift sensitive to processing; batch variability; localized electrode definition and terminal encapsulation often required Electrospinning Polymer/biocomposite solution prep → high-field jetting → controlled collection (alignment/thickness) → solvent removal/drying → crosslinking/stabilization → electrode lamination/embedding + selective encapsulation Breathable epidermal interfaces; wound dressings; porous scaffolds for tissue-contact layers where mass transport is critical Nanofibrous porosity and high surface area improve breathability, wet adhesion, and compliant interpenetration with soft tissue Highly breathable; tunable porosity/thickness; strong mass-transport advantages Mechanical fragility; durability under wet cycling/abrasion; integration of electrodes/interconnects and selective protection becomes the main bottleneck Open in a new tab 4.2. Microfabrication methods Microfabrication remains the primary route for high-resolution patterning, dense interconnects, and controlled multilayer stacks, especially for high-channel-count electrodes and miniaturized sensing/stimulation interfaces. A key advantage for tissue coupling is that microfabrication enables ultrathin laminates and strain-engineered layouts (serpentine/mesh/island–bridge) that reduce effective stiffness and stabilize coupling under deformation ( Fig. 4 ). The main limitations are restricted compatibility with solvent-sensitive or hydrated layers, limited 3D architecture freedom, and the dependence of long-term reliability on encapsulation and connectivity. Fig. 4. Open in a new tab Microfabrication routes positioned by precision and biointerface adaptability. Representative microfabrication strategies for biointerfaces are organized in a qualitative design space defined by patterning precision/integration density and biointerface adaptability. High-precision planar approaches, including film deposition, photolithography and laser patterning, are positioned toward the fine-feature, high-density end, whereas transfer printing, micromolding and soft lithography occupy an intermediate regime that couples microstructural definition with improved compatibility for soft and non-planar substrates. Thermal drawing, wet spinning and electrospinning are grouped as fibrous/porous interface-forming routes with greater adaptability to compliant, curved and hydrated tissues, while micromachining is shown separately as a subtractive route. Representative biointerface outcomes enabled by these method families, including dense electrode arrays, conformal replicated interfaces and fibrous/porous breathable interfaces, are highlighted. Thin-film deposition and barrier construction. PVD-based metals and dielectrics provide high electronic quality and thickness control, but are susceptible to cracking/delamination when residual stress and cyclic strain are not managed 130 – 132 . Conformal dielectric/barrier layers (including vapor-deposited polymer barriers) can improve wet stability and insulation integrity on deformable substrates, but must balance barrier performance with flexibility 131 , 133 . Light-enabled patterning and micromachining. Photopatterning provides reproducible dense features, whereas laser-based patterning enables maskless, rapid iteration and can be effective for structuring soft substrates 119 , 134 , 135 . For wet or bioactive interfaces, practical process flows often separate aggressive patterning steps from hydrated/bioactive layers, shifting emphasis toward multilayer assembly and sealing rather than single-step monolithic fabrication 136 , 137 . Soft lithography and micromolding. Mold-based replication is particularly useful when microstructures must be defined directly in elastomers/hydrogels or when microfluidics are required for biofluid handling 138 – 143 . Compared with planar lithography, these methods provide broader materials compatibility but typically lower feature density and require careful control of sealing and adhesion under wet cycling 138 , 139 . Extrusion-based printing (DIW) and screen printing. DIW enables 3D structures and multi-material deposition using shear-thinning inks, making it well suited for hydrogel/ion-conducting interfaces where 3D geometry and wet compliance drive coupling 144 – 146 . Screen printing offers low-cost large-area manufacturing but coarser resolution 147 , 148 . Both require careful management of curing/drying, layer adhesion, and encapsulation to avoid drift under wet use. In textile manufacturing, screen printing is particularly compatible due to tolerance to rough/porous substrates and high throughput, whereas DIW can deposit viscoelastic inks or hydrogel-like layers onto textile regions for localized soft contact and bioactive interfaces; long-term performance is often governed by adhesion and wash/sweat durability rather than nominal print resolution. Photopolymerization-based 3D printing. Photopolymerization-based 3D printing provides a versatile route to fabricate three-dimensional microarchitectures with high geometric fidelity, making it attractive for bioelectronic interfaces that require precise structural definition and microscale customization. Stereolithography (SLA) builds 3D objects by sequentially curing patterned resin layers, typically through laser scanning along a programmed path, and can produce structures with fine features and smooth surface finishes 149 ( Fig. 5 ). A related strategy forms 3D architectures through repeated 2D photopatterning of thin precursor layers. To improve speed and design flexibility, digital micromirror devices (DMDs) can generate programmable “virtual masks,” avoiding the need for physical mask fabrication 150 , 151 . Fig. 5. Open in a new tab Printing routes positioned by tissue adaptability and multifunctional integration. Representative printing-based fabrication strategies for biointerfaces are organized in a qualitative design space defined by surface/in vivo adaptability and soft-material/3D multifunctional integration. Screen, inkjet and aerosol jet printing are positioned as planar and scalable routes for large-area soft patterning, whereas stereolithography, two-photon polymerization and direct ink writing are placed in the regime of structured three-dimensional fabrication with enhanced architectural and material versatility. In vivo light printing and in vivo sound printing are located toward the high-adaptability end to reflect their ability to form or cure constructs directly on or within living tissues. Representative outcomes, including large-area soft patterning, 3D multimaterial interfaces and in situ tissue-integrated constructs, are indicated to link the fabrication routes to biointerface function. This schematic provides a conceptual framework for comparing printing strategies in biointerface design. These methods, however, often require chemical development to remove uncured regions, which can become challenging for nanocomposites or solvent-sensitive formulations. Advanced variants such as digital light processing (DLP) and continuous liquid interface production (CLIP) further expand throughput and processing flexibility 152 – 155 . Yet, a central limitation for tissue-interfacing applications is that many photocurable resins remain mechanically rigid compared with the soft and compliant materials generally preferred at biological interfaces. Thus, integrating elastomers, hydrogels, or shape-memory polymers into these platforms, while preserving structural integrity and long-term stability, remains a key challenge for their broader use in bioelectronic systems 156 , 157 . Two-photon polymerization (2PP) extends photopolymerization into the regime of highly localized, three-dimensional micro/nanofabrication and is increasingly used for miniaturized medical structures. In 2PP, near-simultaneous absorption of two photons within a tightly confined focal volume initiates photochemistry only at the focal point 158 , 159 ( Fig. 5 ). With femtosecond-pulsed lasers, this approach enables true 3D structuring with feature sizes down to the nanoscale 159 , offering particular value where interface topology, microscale guidance, or device miniaturization are critical. A broad range of photo-responsive materials can be processed by 2PP, including acrylate-based polymers, organically modified ceramics, zirconia sol–gel systems, and titanium-containing hybrid materials. However, the trade-off among resolution, build volume, and printing speed still constrains throughput, and continued advances in optical design, materials chemistry, and process control will be needed to translate these capabilities more broadly into bioelectronic applications. DLP/SLA can produce complex architectures and mechanically programmed constructs, while two-2PP offers nanoscale feature control for microstructured components. For tissue interfaces, resin stiffness, residual species, and wet-state stability can constrain direct coupling, motivating post-processing, composite formulations, or hybrid assembly approaches. Photocurable coatings and UV-curable encapsulants are also frequently used as textile-compatible finishing layers for localized protection of interconnects and nodes without fully blocking airflow. Planar printing methods are generally better suited for scalable soft patterning over large areas, whereas DIW, SLA/DLP, and 2PP become more advantageous when three-dimensional structuring, local multifunctionality, or microscale topology are required. In vivo/in situ printing extends this progression further by directly addressing irregular, dynamic, and site-specific tissue environments, where fabrication and implantation must be integrated into a single workflow. 4.4. In vivo / in situ printing With the development of 3D printing technologies, in vivo bioprinting—also termed in situ, intraoperative, or living bioprinting—has emerged as an approach in which bioinks are deposited directly onto tissues under clinically relevant conditions, enabling immediate interaction with living systems 160 . A key distinction from conventional workflows is that printing and implantation are effectively integrated into a single step with minimal invasiveness. This integration allows real-time adaptation to local surface morphology and defect geometry, so that printed constructs can closely conform to complex tissues. Moreover, in situ crosslinking can enhance adhesion to surrounding tissue and reduce contamination or mechanical damage associated with scaffold handling, transplantation, or suturing 160 – 162 . As in vivo printing forms 3D constructs directly inside the body, it imposes stricter requirements on bioinks than ex vivo printing. Ideal bioinks should support rapid in vivo gelation under physiological conditions, match the mechanical compliance of native tissue, provide biocompatibility with controlled biodegradation, maintain strong wet-tissue adhesion, and enable one-step curing without disruptive post-processing 163 . Light-triggered in vivo printing Light-based in vivo printing typically involves delivering a liquid bioink precursor to a defect site, followed by spatially controlled illumination to induce localized gelation or polymerization. A central requirement is sufficient tissue penetration of the excitation light, which limits the utility of ultraviolet activation in many settings 149 , 164 . Accordingly, considerable effort has focused on near-infrared (NIR)-responsive systems with improved soft-tissue penetration ( Fig. 5 ). For example, nanoparticle-assisted initiation has been used to enable GelMA gelation under DMD-projected patterns, producing subcutaneously printed hydrogels with high cell viability and favorable tissue integration 165 . Alternatively, two-photon excitation of coumarin derivatives has enabled high-resolution in vivo printing without conventional photoinitiators 166 . In these systems, multiphoton microscopy can simultaneously image and guide laser scanning, creating a closed-loop process with micrometer-scale precision. Such minimally invasive light-based printing has been demonstrated in mouse skin, muscle, and brain, although penetration depth and printable volume remain key limitations. Ultrasound-triggered in vivo printing Ultrasound offers a complementary trigger for in vivo printing owing to its deep tissue penetration and compatibility with real-time imaging. High-frequency (>2 MHz) acoustic waves can propagate through heterogeneous tissues, and focused ultrasound can localize energy deposition at depths approaching tens of centimetres. These properties have motivated emerging approaches that use ultrasound as a trigger for in vivo 3D printing ( Fig. 5 ). For example, an “acoustic ink”—a mixture of polymers, particles, and chemical initiators—was designed to absorb acoustic energy, rapidly form a gel to suppress flow, and simultaneously enable a thermally triggered acrylate polymerization reaction, thereby achieving localized material formation 167 . Building on this direction, an ultrasound-enabled 3D printing approach capable of creating complex structures beneath centimeter-thick tissue in vivo was developed. This strategy could be integrated with conventional ultrasound imaging to customize implant geometries based on real-time needs. Importantly, acoustic inks can be engineered with additional functions relevant to bioelectronics, such as electrical conductivity, local drug delivery, and wet adhesion. Incorporating conductive nanoparticles enables the printing of soft sensors and biopotential electrodes that can record electrical activity from living tissues, supporting proof-of-concept monitoring of physiological signals such as ECG and EMG 67 . Extrusion-based and droplet-based in vivo printing In addition to light- and ultrasound-triggered strategies, extrusion-based and droplet-based approaches have been explored for in situ printing, particularly on external or surface-accessible tissues. In extrusion bioprinting, bioinks rapidly respond to a prescribed stimulus applied at the nozzle (e.g., pressure-driven extrusion), forming a continuous, smooth filament that is well suited for macroscale deposition and can support its own weight 168 . Photo-crosslinkable gelatin methacryloyl (GelMA) is a representative extrusion-printable ink, and extrusion-based in situ printing has been widely studied for applications such as skin wound closure and superficial tissue repair, where rapid coverage and scaffold formation are priorities. By contrast, droplet-based in situ bioprinting deposits bioinks as discrete droplets at predefined locations, enabling precise, non-contact placement 169 . Droplets are commonly generated at picoliter scales using thermal, piezoelectric, or laser-assisted actuation, which helps preserve cell viability while offering fast printing, relatively high spatial resolution, and flexible patterning 170 . In situ fabrication merges manufacturing and implantation, enabling constructs that conform to site-specific geometries while avoiding handling damage of fragile soft scaffolds. The key advantages include geometry-adaptive conformity, enhanced wet adhesion via in situ curing, and reduced invasiveness. The primary challenges are triggering safety and penetration depth, reproducibility/quality control in vivo, and control of residual/reactive species and degradation behavior. Light-triggered approaches offer high spatial control but are constrained by penetration depth and printable volume. Ultrasound-triggered strategies provide deeper penetration and image-guided placement, enabling volumetric formation at depth; however, they impose stricter requirements on curing consistency, structural fidelity, and quality control. Extrusion- and droplet-based approaches are effective for surface-accessible tissues but require careful management of wet adhesion, curing kinetics, and defect-free deposition under physiological motion 160 – 162 . 4.5. Combination strategies Soft bioelectronic systems are rarely built as single-function devices; instead, they typically comprise multiple functional modules assembled into an united platform. To realize closed-loop operation that couples diagnosis with therapy, a single system often needs to incorporate sensors, control system, Data analysis and storage, power supply, and stimulation units. A fundamental challenge is that these modules frequently rely on distinct material sets and fabrication requirements, making it highly nontrivial to manufacture every component directly on one fully stretchable substrate without sacrificing performance. To address this gap, hybrid (mixed-mode) manufacturing has become an important route for electronics that must conform to complex biological surfaces. From a tissue-integration perspective, the core concept is mechanical decoupling: strain-isolating interconnects and compliant interface layers preserve stable contact while protecting rigid modules from deformation-induced failure. In hybrid integration, rigid, high-performance electronic elements are combined with soft or flexible substrates, enabling a practical balance between long-term wearability/implantability and high-fidelity signal integrity. By decoupling mechanical compliance from electronic functionality, hybrid integration can preserve high-performance electronic function while reducing interfacial strain and improving mechanical compliance, thereby enhancing user comfort and long-term stability—representing a key enabling concept for advanced bioelectronic interfaces. At the manufacturing level, combining micro/nanofabrication with multiple 3D-printing approaches and tissue-engineering techniques enables the construction of 3D electronic architectures directly on or within biological tissues, with robust functionality and stable performance. Beyond process-level integration, system-level assembly is equally critical: multilayer stacking to form vertically integrated architectures 171 – 174 , adhesive-based bonding of discrete components 171 , 175 , 176 , and mechanical interlocking or co-assembly between different substrates 129 can all support efficient integration of multifunctional biointerfaces within a single platform. Equally critical are packaging and encapsulation, which often dominate long-term drift and failure via moisture/ion ingress, insulation degradation, and adhesion loss under wet cycling. For wearables, permeability and sweat management are critical to prevent secretion accumulation and skin irritation, motivating architectures that maintain breathability while protecting interconnects. For implants, barrier performance and adhesion stability under wet cycling dominate long-term reliability, increasing the importance of encapsulation chemistry, multilayer design, connector robustness, and system-level quality control. For textile-form systems in particular, long-term reliability is frequently limited by textile-to-module interconnects and selective encapsulation strategies that protect routing under sweat and laundering while preserving breathability and compliant skin contact. Taken together, these comparisons indicate that fabrication methods for bioelectronic interfaces should be selected according to the dominant tissue-integration requirement rather than according to resolution or process familiarity alone. Dense planar microfabrication is preferable for highly integrated electrode systems; replication and soft printing strategies are more effective for conformal and breathable interfaces; fiber- and porous-forming routes are advantageous when compliance, permeability, and low bending stiffness are prioritized; and in situ printing is uniquely suited to anatomically irregular or dynamically evolving tissue environments. Hybrid integration is therefore increasingly necessary to combine these strengths within clinically relevant multifunctional systems. 5. Applications in sensing and modulation Bioelectronic devices interfacing with biological tissues have rapidly expanded into wearable, implantable, and ingestible/intraluminal formats for monitoring physiological signals and enabling therapeutic modulation 177 . Across these platforms, practical performance is shaped by a common set of considerations: mechanical conformity and strain accommodation at the interface, time-dependent behavior of soft contact materials, surface and bioactive strategies that regulate biofouling and host response, and manufacturable routes for device fabrication, packaging, and integration. In the following, representative systems are discussed under physiological monitoring, multimodal disease management, and closed-loop operation, while distinguishing wearable, implantable, and ingestible implementations where their interfacial environments impose different design constraints. The representative cases are explicitly mapped to interface archetypes, soft material states, and fabrication routes, as summarized in Tables 2 – 4 . Table 2. Representative health monitoring bioelectronic systems. Name Device format Interfacial strategy Material Fabrication techniques Interface implication Practical constraint / coupling condition 1.3-μm elastic conductor 265 Wearable / skin contact Ultrathin + microstructure (microcracked Au on ultrathin elastomer) Rubber (PDMS) Au thin-film deposition on ultrathin PDMS (supported carrier) → controlled microcrack formation → lamination/encapsulation Ultralow bending stiffness + strain-tolerant microcracks stabilize electrode–skin coupling under motion. Crack morphology and ultrathin handling/edge encapsulation largely determine durability during wear. NeuroGrid 266 Implantable (cortical surface) Ultrathin + microstructure (high-density conformal μECoG array) Rubber-like thin solid (ultrathin polymer stack) * Photolithography-defined electrodes/interconnects on ultrathin polymer film → encapsulation with openings → connectorization Conformal cortical contact reduces micromotion-driven impedance changes while enabling high-channel mapping. Wet insulation integrity, biofouling, and lead-out connector strain relief govern chronic stability. PillTrek capsule 267 Ingestible (GI) Microstructure (multisensor array + controlled exposure windows + sealed architecture) Rubber + other (PDMS + rigid capsule) Precision capsule assembly/sealing + patterned sensor module (laser/printing) + exposure-window definition Performance is dictated by sealing and exposure control under transient mucosal apposition and GI fluids. Mucus/peristalsis-induced variability and fouling at exposed sites drive drift and calibration burden. Fiber–chip hybrid 195 Implantable / intraluminal Microstructure (+ mesh locally) (multifunctional fiber architecture) Rubber + other (soft cladding + structural polymer; hybrid * ) Thermal drawing of multimaterial fiber → post-processing exposure sites → module integration + end packaging Fiber geometry enables deep/tortuous access with distributed interfacing and low effective stiffness. Terminal connectorization, moisture barriers, and fatigue at fiber–module transitions dominate reliability. Knitted washable sensor array 198 Wearable Mesh + microstructure (textile network + sensing nodes; selective protection) Rubber-like solid (textile network) * Knitting/weaving functional yarn network → node integration → localized encapsulation/routing protection Breathable large-area contact shifts coupling to garment-scale mechanics and distributed sensing. Motion/contact-pressure variability and wash/sweat durability depend on routing and selective encapsulation. Open in a new tab * Note: * “rubber-like thin solid” denotes compliant polymer films; “hybrid” denotes mixed rigid-soft stacks with localized soft contact layers. Table 4. Representative closed-loop bioelectronic systems. Name Device format Interfacial strategy Material Fabrication techniques Interface implication Practical constraint / coupling condition Soft neuromorphice-skin 272 Wearable closed-loop Ultrathin/mesh (system-level) (distributed sensors + local neuromorphic conditioning; compliant coupling) Rubber (hybrid) (soft elastomeric electronics) Soft electronics integration (distributed sensors + compliant interconnects + encapsulation) Compliant epidermal coupling with local signal conditioning enables low-voltage, on-body feedback operation. Sweat/motion-driven coupling variability and drift set the stability of state estimation and control outputs. Chronic adaptive DBS 209 Implantable closed-loop Microstructure (biomarker sensing + stimulation gating; clinically tuned control policy) Other (hermetic implant) Medical-grade lead + hermetic implant manufacturing with validated packaging/connectors Biomarker-informed gating links chronic sensing to stimulation dosing for adaptive neuromodulation in patients. Long-term biomarker stability, sensing drift, and calibration/latency constraints govern sustained closed-loop benefit. Gastric-retentive robotic capsule 269 Ingestible closed-loop Microstructure (retention + sensor/trigger + actuator for responsive delivery) Other + liquid payload Precision capsule assembly + sealing (retention/actuation module integration + reservoir packaging) Retention plus responsive actuation enables closed-loop intervention in the GI tract without permanent implantation. Safety of long residence, retention control, and power/telemetry robustness dominate practical deployment and regulation. Bidirectional fiber platform 195 Implantable / intraluminal Microstructure (+ mesh locally) (distributed sensing/stimulation nodes compatible with feedback control) Rubber + other (hybrid) Thermal drawing of multifunctional fiber + module integration + end packaging/connectorization Distributed fiber nodes support bidirectional interfacing suitable for feedback control in deep tissue and tortuous anatomy. End-connectorization, wet insulation at exposed sites, and fatigue at fiber–module transitions limit longterm closed-loop stability. Embroidered liquid-metale-textiles 273 Wearable closed-loop enabler Mesh (garment-integrated power/comm backbone for distributed loops) Liquid (Galinstan) + rubberlike solid (textile) Liquid-metal fiber fabrication (infusion) + digital embroidery + selective encapsulation/routing Garment-scale power/communication backbones enable distributed sensing/actuation loops without rigid wiring. Encapsulation must balance washability and breathability; stitched interconnect fatigue and EMI sensitivity affect robustness. Open in a new tab 5.1. Physiological signal monitoring Tissue-interfaced sensing enables continuous acquisition of physiological signals directly from the body, supporting real-time health monitoring and disease management. Multiple sensing modalities, including electrophysiological ( Fig. 6a – c ) 178 , biochemical ( Fig. 6d – f ) 179 – 181 and biophysical ( Fig. 6g – i ) approaches 19 , 182 , 183 , can be implemented through appropriately engineered tissue–device interfaces. In practice, performance is frequently limited by interface coupling stability, biofluid handling and fouling, and biological remodeling over time. Fig. 6. Open in a new tab Representative tissue-interfaced bioelectronic platforms for physiological monitoring. Blue, green, and orange markers denote wearable, implantable, and ingestible platforms, respectively; dual markers indicate systems that span more than one form factor. (a–c) Electrophysiological monitoring. (a) Ultrathin soft sensor platform for epidermal and implantable ECG monitoring, figure was reproduced with permission from ref. 178 , Copyright 2022 Springer Nature. (b) Highly compliant cortical interface (NeuroGrid) for stable LFP recording, figure was reproduced with permission from ref. 189 , Copyright 2015 Springer Nature. (c) Machine-knitted, washable sensor-array textile for epidermal pulse monitoring, figure was reproduced with permission from ref. 198 , Copyright 2015 American Association for the Advancement of Science. (d–f) Biochemical sensing. (d) Skin-interfaced wearable aptamer sensor for hormone monitoring in sweat, figure was reproduced with permission from ref. 180 , Copyright 2024 Springer Nature. (e) Ingestible smart capsule for longitudinal gastrointestinal biochemical profiling, figure was reproduced with permission from ref. 181 , Copyright 2025 Springer Nature. (f) Microneedle/interstitial-fluid platform with regenerable protein sensors for continuous inflammation monitoring, figure was reproduced with permission from ref. 179 , Copyright 2024 American Association for the Advancement of Science. (g–i) Biophysical sensing. (g) Conformal piezoelectric actuator-sensor array for near-surface tissue viscoelasticity measurement, figure was reproduced with permission from ref. 95 , Copyright 2015 Springer Nature. (h) Injectable bioresorbable ultrasonic sensor for deep-tissue multiparameter monitoring, figure was reproduced with permission from ref. 182 , Copyright 2024 Springer Nature. (i) Implantable thermal sensor system for organ-level inflammatory assessment. LFP, local field potential; ECG, electrocardiogram, figure was reproduced with permission from ref. 19 , Copyright 2023 American Association for the Advancement of Science. 5.1.1. Wearable monitoring (epidermal interfaces) Electrophysiological signals Electrophysiological sensors capture bioelectric potential variations from electrically active organs and tissues, most commonly the brain, heart, and skeletal muscle, yielding signals such as EEG, ECG, EMG, and EOG. In practice, these measurements typically require intimate electrode tissue contact to maximize coupling efficiency and sensitivity. Because the electrode tissue interface impedance directly shapes signal quality (noise floor, bandwidth, and motion artifacts), engineering a high-quality, conformal biointerface is often as important as the front-end electronics. A representative example is an ultrathin, breathable elastic conductor built on a thin elastomeric substrate with a microcracked metal film, achieving a total thickness on the order of ~1.3 μm. Its low bending stiffness enables nearly seamless lamination, supporting stable ECG acquisition and detection of subtle mechanical signatures such as pulse-wave-related signals 178 ( Fig. 6a ). This performance arises from the combination of an ultralow bending-stiffness architecture and a stretch-tolerant microcracked conductor on an elastomeric substrate, together with scalable thin-film patterning and robust lamination/encapsulation that preserve sub-micron flexibility during handling and wear. At the non-invasive end of the spectrum, a wrist-worn surface-EMG neuromotor interface demonstrates that robust motor-intent decoding can be achieved with wearable electrophysiology, balancing performance with usability for human–computer interaction. In such systems, practical stability is often achieved through strain-isolating layouts that decouple rigid electronics from compliant contact regions, supported by soft band mechanics and reliable packaging/assembly; the resulting robustness and user-friendliness come with trade-offs in ultrathin lamination and a higher dependence on placement and contact pressure. A complementary, clinically oriented example is the class of skin-interfaced, wireless physiological monitoring systems deployed in neonatal and pediatric intensive-care settings 184 . In this regime, translational constraints—such as gentle skin coupling, reliable wireless operation, and long-duration stability under clinical handling—become as important as peak signal quality, and robust packaging and compatibility with clinical workflows strongly shape practical adoption. Biochemical biomarkers Biochemical markers, such as pH, glucose, lactate, electrolytes, and hormones, provide critical information on physiological homeostasis and disease states. These targets can be sampled from multiple biofluids (blood, sweat, interstitial fluid, saliva, and gastrointestinal fluids), but continuous operation, sensing reliability, long-term stability, and user comfort remain persistent constraints for many biochemical biointerfaces. A skin-interfaced wearable aptamer platform based on a target-triggered strand-displacement mechanism provides a concrete pathway toward non-invasive hormone tracking (e.g., estradiol) through in situ sweat analysis, supported by on-body microfluidics for time-resolved sampling 180 ( Fig. 6d ). Here, controllable sampling and transport at the skin–biofluid interface, together with antifouling surface chemistry that preserves sensor responsiveness in complex sweat matrices, are as important as the recognition element itself; manufacturable integration typically relies on patterned electrodes combined with microfluidic structuring and dependable multilayer sealing to limit leakage, evaporation, and drift. In parallel, interstitial-fluid microneedle arrays provide an alternative route that reduces sampling variability relative to sweat while remaining minimally invasive. Tehrani et al. reported a fully integrated wearable microneedle array enabling wireless, continuous real-time sensing of metabolite pairs in human volunteers during daily activities 185 . Such systems highlight a complementary design space in which analytical robustness depends on stable dermal coupling, biofouling control at sensing interfaces, and system-level calibration and quality control for long-duration operation. Biophysical signals Bioelectronic interfaces increasingly target biophysical readouts, including mechanical forces, temperature, and optical or acoustic signatures. Strain and pressure sensors quantify tissue deformation and force transmission, respectively. A conformal piezoelectric actuator–sensor network built from ultrathin inorganic piezoelectric elements (e.g., PZT nanoribbons) enables rapid, quantitative assessment of near-surface tissue viscoelasticity with intimate contact to soft, textured biological surfaces 95 ( Fig. 6g ). The quantitative mapping capability in this case is governed by efficient strain transfer enabled by an ultrathin, compliant network architecture and soft encapsulation/adhesion, together with fabrication routes that preserve alignment, encapsulation integrity, and repeatable contact on textured skin. Finally, bioacoustic/mechano-acoustic sensing has become increasingly attractive for non-invasive monitoring of deep or otherwise hard-to-access physiology. Many routine biological activities (e.g., valve motion, gastrointestinal peristalsis, vocal fold vibration) generate mechanical waves with rich spectral content; soft, skin-compatible acoustic sensors can capture these signatures for continuous health monitoring and human-machine interfacing 186 – 188 , practical performance is governed by interfacial coupling stability, biofluid interactions, material time-dependent behavior, and fabrication/packaging constraints. Across wearable systems, interface stability under motion, sweat-induced biofluid interactions, and packaging durability consistently limit long-term signal reliability. 5.1.2. Implantable monitoring (surface and deep-tissue interfaces) Implantable interfaces must maintain stable coupling in wet, mechanically dynamic tissue while operating under biofouling and foreign-body response. Long-term drift is often driven by micromotion-induced tissue injury, progressive encapsulation, and degradation of insulation or adhesion, motivating co-design of compliant mechanics, surface regulation, and robust packaging. For neural recordings, an ultra-compliant and scalable cortical interface (NeuroGrid) illustrates how organic, highly conformal arrays can follow the brain’s curved topology and deliver stable, high-SNR field-potential recordings in clinical contexts such as epilepsy monitoring 189 ( Fig. 6b ). In this setting, topology-matched compliance reduces micromotion and stabilizes impedance, while long-term performance is frequently shaped by surface-mediated biofouling and tissue responses; accordingly, microfabrication-based electrode patterning and polymer encapsulation must be coupled with packaging/connectorization strategies that preserve flexibility and insulation reliability during chronic use. Beyond planar cortical arrays, deep-tissue monitoring is also beginning to adopt fiber-like implant architectures to expand spatial coverage while limiting invasiveness. In this context, fiber-form interfaces offer a complementary route to distributed access, organ-specific adaptability, and reduced mechanical footprint, although their broader interfacial advantages and representative implementations are discussed in more detail in Section 5.1.4 . For higher-molecular-weight targets, microneedle-based interstitial fluid platforms combined with regenerable (“active-reset”) protein sensors have demonstrated real-time tracking of protein biomarkers in vivo, representing an important step for translating biochemical sensing beyond small molecules 179 ( Fig. 6f ). This approach leverages microstructured access to interstitial fluid with minimal invasiveness, but sustained accuracy depends on limiting fouling and enabling sensor regeneration; practical translation is therefore tightly coupled to microneedle manufacturing yield, tip robustness, integration of microfluidic/electrode components, and reproducible insertion performance. For deep-tissue monitoring, an injectable, bioresorbable metastructured hydrogel (“metagel”) ultrasonic sensor demonstrates wireless multiparameter readout (including intracranial pressure, temperature, pH, and flow-related signals) by translating volumetric responses into ultrasound-detectable spectral shifts, with reported sensing at depth and degradation over weeks, highlighting a minimally invasive route to chronic monitoring without wired leads 182 ( Fig. 6h ). Importantly, this sensor is designed to be embedded within tissue: viscoelastic stress relaxation supports stable embedded coupling, while the internal periodic architecture encodes an ultrasound-readable spectral response; reproducible casting/packaging of the microstructured channels is therefore central to sensing fidelity, stability under biological remodeling, and programmed biodegradation. Temperature monitoring is also clinically meaningful because localized thermal changes can serve as early indicators of inflammation or infection. An implantable bioelectronic system that continuously tracks local organ temperature and thermal conductivity on the kidney surface illustrates how biophysical sensing can support early detection of inflammatory processes associated with transplant rejection in animal models 19 ( Fig. 6i ). Here, long-term reliability is enabled by compliant, strain-tolerant layouts that accommodate organ motion while maintaining wet contact, combined with durable soft encapsulation and system-level integration that limits signal drift and preserves insulation integrity. Across implantable systems, micromotion, biofouling, and encapsulation-driven drift consistently emerge as dominant constraints on long-term performance. 5.1.3. Ingestible / intraluminal monitoring (GI interfaces) Ingestible biointerfaces extend biochemical sensing to internal organs, but the dominant constraints differ from chronic implants: operation occurs under variable pH, enzymes, mucus, and peristaltic forces, with transient and intermittent tissue apposition. As a result, packaging robustness, exposure-window design, and mucus-mediated fouling control often dominate sensing fidelity. The PillTrek capsule (~7 mm diameter, 25 mm length) integrates miniaturized wireless electrochemical instrumentation with interchangeable sensors to profile multiple gut parameters (e.g., pH, temperature, neurotransmitters, and metabolites), enabling longitudinal tracking of dynamic monitoring of diet- and health-related biochemical dynamics 181 ( Fig. 6e ). In this intraluminal setting, measurement stability is governed less by long-term fibrosis mitigation and more by transient mucosal apposition, mucus fouling, and GI-robust sealing; accordingly, capsule architecture, exposure control, and anti-fouling surface chemistry become primary determinants of practical performance. Across ingestible systems, biofluid variability, mucus-mediated fouling, and sealing reliability dominate sensing stability and device performance. 5.1.4. Fiber and textile-form biointerfaces Beyond planar laminates and patches, one-dimensional (1D) fiber bioelectronics and electronic textiles provide distinct form factors for biointerfaces. As noted above for deep-tissue monitoring, fiber-shaped devices can complement conventional planar or shank-based implants by combining small diameters and low bending stiffness with the ability to navigate spatially constrained or tortuous anatomy, thereby enabling distributed sensing and actuation with minimal added mechanical burden. Textile-form interfaces, by contrast, leverage the hierarchical architecture of fibers–yarns–fabrics to achieve large-area coverage with high breathability and wearing comfort, which can mitigate limitations of skin-adhering patches such as restricted airflow and prolonged reliance on adhesives. In both platforms, practical performance is jointly determined by mechanical coupling under motion, wet-state stability and biofouling control, and manufacturable integration of interconnects and packaging. Fiber bioelectronics: deep-tissue access and procedure-compatible integration A landmark example of multifunctional deep-tissue fiber probes is the class of thermally drawn multimaterial fibers integrating electrical recording/stimulation with optical and microfluidic functions. Canales et al. reported a multifunctional fiber platform that combines multiple modalities within a single compliant probe 190 . Beyond the form factor advantage, this approach illustrates how continuous, scalable manufacturing (preform assembly and thermal drawing) can embed multifunctionality along the fiber axis, while tissue compatibility is dominated by stable mechanical coupling (low bending stiffness), reliable insulation/encapsulation in wet environments, and robust end-connectorization, which often becomes the limiting factor for chronic reliability. A complementary landmark demonstration is the stretchable multimodal fiber probe for spinal cord interfacing reported by Lu et al. 191 . In this work, a mechanically compliant fiber integrates an optical waveguide with a stretch-tolerant conductive mesh, enabling recording of sensory-evoked and optically evoked potentials after implantation and maintaining coupling under deformation. This example highlights a recurring design constraint for moving, deformable tissues: long-term function depends not only on compliance but also on wet-state stability of conductive coatings/meshes, strain-tolerant encapsulation, and fatigue-resistant interconnect transitions at the fiber ends, where mechanical mismatch and moisture ingress can dominate failure. Beyond deep neural access, fiber biointerfaces can also be aligned with standard clinical workflows. Kim et al. introduced bioelectronic sutures with electrochemical pH sensing for longterm monitoring of wound healing progress 192 , demonstrating that thread-like interfaces can provide intimate coupling at surgical sites with minimal added bulk. In such suture-like systems, translation is often governed by selective exposure versus encapsulation (protecting interconnects while keeping sensing regions accessible), fouling and calibration drift in exudate-rich environments, and sterilization/storage constraints. Related wearable implementations also exist, such as fiber-based electrochemical biosensors incorporated into garments for sweat glucose monitoring 193 , underscoring that fiber architectures can bridge implantable and wearable contexts, while again placing emphasis on coating adhesion, abrasion resistance, and stable electrical routing under repeated bending and wetting. Recent fiber-based systems further broaden the application scope by emphasizing spatial coverage, deployment flexibility, and organ-specific interfacing. A representative example is the spatially expandable fiber-based probe guided by a helical scaffold reported by Jiang et al. 194 , in which the implanted fiber can access multiple deep-brain regions from a single insertion pathway. This strategy illustrates a distinctive advantage of fiber architectures over conventional rigid shanks: rather than merely minimizing footprint, they can actively reconfigure the sampling geometry in three dimensions after implantation, thereby improving distributed access while preserving a relatively compliant mechanical interface. At the same time, such systems make clear that practical performance depends not only on the fiber body itself, but also on controlled deployment, stable anchoring, and reliable interconnect transitions at the externalized end. An additional step toward organ-specific multifunctionality is illustrated by the microelectronic fiber–chip hybrid platform reported by Sahasrabudhe et al. 195 , which combines thermally drawn fiber backbones with chip-level electronic functionality to enable wireless interfacing with both gut and brain circuits. This work highlights an important maturation of the fiber biointerface concept: fibers are no longer used only as passive compliant carriers or simple multimodal conduits, but increasingly serve as system-level platforms that integrate sensing, stimulation, wireless communication, and organ-adapted mechanics within a single slender architecture. In this regime, long-term robustness is governed by the full device chain, including wet-state encapsulation, connectorization or chip–fiber integration, and resistance to mechanical fatigue and moisture ingress under chronic operation. Together, these examples show that fiber architectures can extend sensing and modulation beyond planar footprints not only by reducing invasiveness, but also by enabling three-dimensional coverage, anatomical adaptability, and multifunctional integration. They also reinforce a recurring translational constraint across fiber biointerfaces: chronic reliability is often limited less by the nominal compliance of the fiber shaft than by packaging, terminal integration, and long-term electrical and mechanical stability in wet biological environments. Electronic textiles: breathable, large-area epidermal monitoring and integrated functions A landmark direction in textile biochemical monitoring is the woven electrochemical fabric reported by Wang et al. 196 , which weaves sensing fibers into a textile to enable real-time health monitoring. This class of platforms illustrates how textile manufacturing (weaving/knitting) naturally supports distributed, breathable interfaces at garment scale, but also shifts key bottlenecks toward interconnect routing, contact-pressure variability, and motion artifacts, motivating textile-specific contact engineering and robust packaging strategies for wash/sweat exposure. For textile electrophysiology, Wang et al. demonstrated high-performance textile electrodes via laser writing of Janus graphene/Kevlar textiles 197 , enabling high-quality ECG acquisition while maintaining fabric flexibility. This example highlights the value of textile-compatible patterning routes (e.g., direct laser writing/printing) for creating conformal electrodes without conventional gels; however, long-term performance remains sensitive to skin–textile contact mechanics, sweat accumulation, and wash durability, making selective encapsulation (protecting traces while preserving breathability at the contact region) a practical determinant of real-world stability. For biophysical monitoring, Fan et al. reported a machine-knitted, washable sensor-array textile for precise epidermal physiological signal monitoring 198 , demonstrating that textile manufacturing can deliver scalable, repeatable arrays while preserving wear comfort and breathability. In general, electronic textiles shift interface bottlenecks from single-device coupling toward interconnect reliability, washability, and long-term mechanical durability, where the best-performing systems typically combine textile-level mechanical compliance with localized reinforcement/encapsulation to protect functional nodes without sacrificing airflow and comfort. Comparative considerations across monitoring form factors. Across physiological monitoring modalities, wearable patches and textiles prioritize motion robustness, breathability, and user comfort, with dominant failure modes in contact variability, sweat exposure, and interconnect durability. Minimally invasive microneedle interfaces reduce sampling variability relative to sweat-based assays but introduce additional constraints in calibration drift and long-term biofouling control at sensing sites. Implantable platforms face chronic drift driven by micromotion and host response, making encapsulation integrity and packaging/connectorization decisive for long-duration stability. Ingestible interfaces shift the bottleneck toward sealing and transient mucosal apposition under mucus-rich environments. Fiber-form architectures provide a complementary path to distributed deep-tissue access and organ-specific interfacing, but again place a premium on end-terminals, encapsulation, and wet-state mechanical/electrical stability. Across fiber and textile platforms, interconnect reliability, wet-state stability, and mechanical durability under repeated deformation are key limiting factors. 5.2. Multimodal disease management Real-time disease intervention and management constitute another major application area of bioelectronics interfaces. In particular, the effective integration of strategies based on diverse stimuli, including pharmacological treatment 199 , electrotherapy 200 , 201 , photostimulation 202 , thermal stimulation 203 , and acoustic stimulation 204 , with bioelectronics devices offers reliable, safe, and scalable possibilities for treating various diseases, especially chronic conditions ( Fig. 7 ). Across modalities, interfacial engineering determines whether actuation is spatially confined, repeatable, and safe, while fabrication and packaging determine whether these functions can be reliably integrated into a clinically deployable form factor. Fig. 7. Open in a new tab Representative tissue-interfaced bioelectronic platforms for therapeutic intervention. Blue, green, and orange markers indicate wearable, implantable, and ingestible systems, respectively. (a–c) Drug-based and intracellular therapeutic platforms: (a) wireless wearable system for wound sensing, electrically triggered drug release, and stimulation-assisted healing, figure was reproduced with permission from ref. 205 , Copyright 2023 American Association for the Advancement of Science. (b) electronic dura mater (“e-dura”) for localized pharmacological modulation and spinal neuromodulation, figure was reproduced with permission from ref. 208 , Copyright 2015 American Association for the Advancement of Science. (c) flexible nanofluidic intracellular delivery patch for electroporation-assisted molecular transport within internal organs, figure was reproduced with permission from ref. 199 , Copyright 2025 Springer Nature. (d–f) Electrical stimulation and electrotherapy: (d) millimetre-scale bioresorbable implant for temporary cardiac pacing, figure was reproduced with permission from ref. 201 , Copyright 2025 Springer Nature. (e) self-powered biodegradable electronic bandage for accelerated intestinal wound repair, figure was reproduced with permission from ref. 210 , Copyright 2024 Springer Nature. (f) conductive hydrogel-based elastic microelectronic array for low-voltage neuromodulation, figure was reproduced with permission from ref. 211 , Copyright 2019 Springer Nature. (g–i) Optical, acoustic, and thermal modalities: (g) multifunctional microelectronic fiber for optogenetic modulation in deep-brain and gastrointestinal applications, figure was reproduced with permission from ref. 202 , Copyright 2024 Springer Nature. (h) wearable flexible ultrasound microneedle patch (wf-UMP) for acoustic-enhanced cancer therapy, figure was reproduced with permission from ref. 204 , Copyright 2025 Springer Nature. (i) flexible wireless thermal modulation system integrating heating, temperature sensing, and drug delivery for tumor treatment, figure was reproduced with permission from ref. 203 , Copyright 2019 Springer Nature. 5.2.1. Wearable therapeutic platforms Soft bioelectronic interfaces provide platforms for localized drug delivery, enabling spatially confined and dose-controllable release at the target site. This localized paradigm can maximize drug exposure in the lesion region while minimizing off-target distribution and systemic side effects. Importantly, drug-delivery functionality can be coupled with in situ sensing of local physiological signals to create feedback-regulated therapeutic systems. A fully integrated, wireless wearable system combines multimodal wound monitoring with electrically triggered release of anti-inflammatory/antimicrobial agents, alongside exogenous electrical stimulation to promote tissue regeneration 205 ( Fig. 7a ). Such wearable therapeutic interfaces typically rely on conformal, strain-tolerant architectures and wet-compatible soft layers that sustain adhesion and ionic/electrical coupling in exudate-rich environments; practical performance is therefore constrained by multilayer integration, sealing reliability, and surface/biofluid fouling that can induce sensor drift and dosing variability. A closely related landmark demonstration is a wireless, closed-loop smart bandage integrating sensors and stimulators for advanced wound care 206 . This class of systems emphasizes that, in addition to sensing accuracy and stimulation efficacy, practical wound-therapy deployment depends strongly on wet-adhesion management, on-demand attachment and detachment to avoid secondary tissue damage, and stable operation under exudate-rich biofouling conditions—constraints that are rarely captured by isolated benchtop metrics. Wearable optical therapeutics provide another complementary path. For example, a wearable and stretchable dual-wavelength LED device has been reported for at-home care of chronic infected wounds 207 , highlighting that therapeutic outcomes depend not only on optical dose delivery but also on conformal contact, thermal management, and packaging reliability during repeated use in home-care conditions. Ultrasound-triggered interventions are also rapidly evolving. An integrated wearable flexible ultrasound microneedle patch combines a conformal lead-free transducer array with a dissolvable microneedle drug-delivery module; ultrasound activation promotes drug release and reactive oxygen species generation for effective tumor therapy, offering a portable and minimally invasive acoustic platform for cancer treatment 204 ( Fig. 7h ). In this case, stable energy delivery depends on consistent acoustic coupling and reliable microneedle–tissue contact, while manufacturability and shelf stability depend on uniform microneedle fabrication, transducer integration, and the longterm stability of soft coupling/adhesive layers during storage and use. Across wearable therapeutic systems, stable interface coupling, controlled delivery, and packaging reliability are critical for achieving reproducible and safe treatment outcomes. 5.2.2. Implantable and bioresorbable therapeutic platforms Implantable neural interfaces have similarly incorporated localized pharmacological actuation. The electronic dura mater (“e-dura”) integrates electrophysiological recording, signal routing, and local delivery modules, enabling electrochemical neuromodulation of spinal circuits and supporting motor-function restoration after severe spinal cord injury 208 ( Fig. 7b ). This multifunctional implant couples dura-matched compliance and conformal contact with integrated delivery channels, such that long-term efficacy depends on micromotion suppression, controlled tissue response, and reliable multi-material integration; in practice, connect and packaging frequently dominate translational robustness. Emerging intracellular delivery concepts further expand the scope of bioelectronic-tissue interactions. A battery-free and chipless flexible nanofluidic intracellular delivery patch (NanoFLUID) has been developed for efficient payload transport into cells within internal organs; the nanopore–microchannel–microelectrode architecture enables safe electroporation-assisted transmembrane delivery, boosting transport kinetics by ~10 5 compared with passive diffusion, and illustrates the potential of stable, cell-level bioelectronic interfacing 199 ( Fig. 7c ). Here, delivery efficiency is tightly coupled to the integrity of micro/nanostructured pathways, stable electrode–tissue coupling, and robust bonding/encapsulation; yield, sealing reliability, and structural verification become practical bottlenecks for translation. Electrical stimulation is widely used to modulate neural and muscular activity for suppressing pathological excitability, relieving pain, or accelerating tissue repair. Flexible and soft stimulation platforms are particularly attractive because their compliant mechanics and conformal contact can reduce interfacial impedance and improve charge-injection efficiency, thereby supporting effective and potentially long-term interventions. A millimetre-scale bioresorbable optoelectronic system that combines wireless power with optical control to deliver programmable stimulation, demonstrating a temporary pacing strategy while minimizing patient burden through fully resorbable materials and minimally invasive deployment 201 ( Fig. 7d ). For such transient implants, the key design constraint shifts toward predictable functional lifetime and benign degradation products, placing a premium on materials selection and encapsulation strategies that align device fate with therapeutic time windows. Recent bioresorbable stimulation systems further extend this concept to long-duration regeneration. A bioresorbable, wireless dual-site stimulator has been reported for peripheral nerve regeneration 209 , illustrating how multi-site stimulation and operation over months can be achieved while avoiding device removal. This direction highlights the practical need to co-design functional lifetime, encapsulation/barrier stability, and controlled resorption to match biological healing timelines. Soft and biodegradable electrotherapeutic devices have also shown substantial promise in regenerative medicine. A self-powered electronic bandage made from soft, bioresorbable materials has been reported to accelerate intestinal wound healing by integrating two complementary stimulation modes including pulsed stimulation to facilitate epithelial cell transfection and direct-current stimulation to enhance pro-healing factor secretion, thereby improving closure and regeneration in vivo 210 ( Fig. 7e ). Here, stable wet adhesion and mechanical compliance must be balanced with controlled degradation, and manufacturing routes must integrate porous scaffolds, patterned electrodes, and hydrogel-like coupling layers without compromising robustness. In addition, conductive hydrogel-based elastic microelectronic arrays, with Young’s modulus in the kPa range and thickness down to ~20 μm, further demonstrate stable in vivo electrical performance and therapeutically relevant current densities at low operating voltages, highlighting the advantages of tissue-like mechanics for efficient neuromodulation 211 ( Fig. 7f ). In these systems, soft material constitutive behavior improves charge transfer and reduces irritation, but long-term aqueous stability, encapsulation integrity, and batch-to-batch reproducibility become central to practical deployment. Thermal modulation provides yet another route to spatiotemporally controlled therapy. A flexible, adhesive, biodegradable, and wirelessly controlled intracranial device has been reported to integrate a wireless heater, a temperature sensor, and a drug-loaded patch; radiofrequency magnetic-field activation enables mild thermal actuation that enhances drug release and penetration depth, yielding therapeutic benefit in a mouse glioblastoma model 203 ( Fig. 7i ). In this setting, therapeutic reliability depends on conformal wet contact in the resection cavity, controlled thermal dosing, and reproducible drug-release kinetics, while multilayer integration and sterilization-compatible packaging impose additional translational constraints. Across implantable therapeutic systems, interface stability, controlled actuation, and long-term material and packaging reliability govern clinical performance. 5.2.3. Intraluminal / GI therapeutic interfaces Soft bioelectronics are increasingly integrated with light, acoustic, and thermal modalities to expand the therapeutic design space. Multifunctional polymer-fiber bioelectronic interfaces that embed multiple solid-state microelectronic elements can access both deep-brain regions and the tortuous gastrointestinal tract. When paired with wireless control electronics, such interfaces can support programmable optogenetic stimulation of dopaminergic neurons to evoke reward-related behaviors, and intraluminal stimulation of vagal-afferent pathways to probe brain–viscera communication 202 ( Fig. 7g ). Intraluminal translation is governed by navigability in tortuous anatomy, resistance to GI fouling, and robust connectorization and packaging, which together determine whether stimulation can be delivered reliably under peristalsis and mucosal contact variability. Fiber-form architectures provide a complementary route for minimally invasive deep-tissue and intraluminal interfacing, with additional representative examples discussed in Section 5.1.4 . Across therapeutic systems, stable interface coupling, controlled energy or drug delivery, and reliable packaging are critical for achieving safe and repeatable interventions. 5.3. Closed-loop strategies Recent advances have enabled biointerfaces that integrate sensing, computation, and actuation into closed-loop systems, advancing the goal of personalized and adaptive therapy 212 . Fig. 8 summarizes the key components and operating principles of closed-loop bioelectronic platforms. Closed-loop performance is fundamentally limited by interface stability: sensor drift degrades state estimation, while unstable actuation interfaces lead to variable dosing/stimulation. Fig. 8. Open in a new tab Architecture of integrated closed-loop soft bioelectronic systems. Schematic representation of a closed-loop bioelectronic platform that integrates sensing, computation, power management, and therapeutic actuation. Physiological signals acquired by conformal sensing units are transmitted to embedded control electronics for signal processing, data storage, and decision-making. Based on predefined algorithms or adaptive models, stimulation or drug-delivery units are activated to provide targeted intervention. Wireless communication modules enable external data transfer and system monitoring, while power sources (for example, batteries or wireless energy transfer modules) support continuous operation. Clinical neuromodulation provides a clear illustration of how closed-loop control can improve real-world outcomes when reliable biomarkers and stable interfaces are available. For example, chronic adaptive deep brain stimulation has been evaluated against conventional stimulation in Parkinson’s disease 209 , demonstrating how biomarker-driven feedback control can reduce symptom burden while highlighting practical requirements in long-term sensing stability, algorithm robustness, and clinically acceptable control policies. Closed-loop paradigms also extend to rehabilitation and seizure management. Closed-loop vagus nerve stimulation paired with rehabilitation has been reported to aid recovery from spinal cord injury 213 , emphasizing that closed-loop triggers can be tied to behavioral success signals and that therapeutic benefit depends on stable sensing, low-latency triggering, and reliable stimulation delivery. In epilepsy models, closed-loop electrical stimulation has been used to prevent disease progression and long-term impairment 214 , illustrating a distinct closed-loop logic in which detection of pathological events gates stimulation to suppress downstream network responses. At the device level, closed-loop capability increasingly relies on miniaturized, wireless system integration that reduces tethers and latencies. A wireless and battery-less implant has been reported for autonomous recording of EEG, EMG, and temperature, together with closed-loop neuromodulation via optogenetics and pharmacology in freely moving small animals 215 , illustrating how on-device computation and untethered operation can enable more naturalistic closed-loop studies while raising stringent requirements for packaging, telemetry reliability, and long-term wet stability. In prosthetics, closed-loop concepts restore motor function by incorporating sensory feedback. A monolithically integrated, soft, low-voltage “electronic skin” (e-skin) platform operates without rigid electronic components while reproducing several core features of biological somatosensation. This system integrates pressure and temperature sensors with neuromorphic circuitry capable of conditioning signals into spike-like pulse trains that directly drive closed-loop actuation, thereby emulating sensory–motor feedback pathways. By addressing long-standing challenges in system-level integration, this work provides a practical route toward e-skins with advanced tactile feedback capabilities 216 . In practice, stable closed-loop performance requires repeatable skin coupling under motion and sweat exposure, together with scalable integration routes that minimize rigid components and preserve mechanical compliance. Closed-loop operation has also been demonstrated in implantable optoelectronic biointerfaces. Miniaturized bio-optoelectronic implants based on microscale inorganic LEDs (μ-ILEDs) deliver optical stimulation to activate opsins. while flexible biophysical sensors continuously track organ function. Embedded control and data-analysis routines detect pathological patterns and automatically trigger intervention. In a representative demonstration, algorithm-guided closed-loop optogenetic modulation of bladder sensory afferents restored normal bladder function 217 . For such implantable closed-loop systems, long-term robustness depends on limiting biofouling and host response to reduce drift, maintaining stable electromechanical coupling under micromotion, and ensuring reliable packaging, sterilization compatibility, and system-level integration. Comparative considerations across closed-loop paradigms. Closed-loop platforms span a spectrum from wearable neuromorphic skins to fully implantable neurostimulators and wireless autonomous implants. Across these systems, key differentiators include the choice of feedback variable (electrophysiological biomarkers, biophysical signals, or behavioral success triggers), permissible latency and computational complexity, and the stability of sensing/actuation interfaces under real-world motion and wet exposure. Clinical closed-loop neuromodulation prioritizes long-term safety, robust control policies, and clinically meaningful endpoints, whereas preclinical closed-loop implants offer greater flexibility in modality integration but face stringent constraints in power, telemetry, and packaging reliability. In all cases, long-term drift at the tissue interface remains a central limiting factor that couples device design and algorithm performance. Across closed-loop systems, long-term interface stability remains the central limiting factor, as drift in sensing or actuation directly degrades control accuracy and therapeutic efficacy. 6. Strategies for biocompatibility optimization Interactions at the bioelectronic–tissue interface largely determine the biocompatibility, performance, and operational lifetime of bioelectronic devices. When interfaces with pronounced mechanical or chemical mismatches are chronically implanted, or maintained in long-term contact with target tissues, adverse events can arise across acute and chronic stages. In neural implantation, for instance, acute complications may include elevated intracranial pressure, hemorrhage, and disruption of the blood-brain barrier, followed by immune-driven recruitment and accumulation of nonfunctional glial cells and astrocytes that progressively compromise the brain-electrode coupling and signal reliability 218 , 219 . In vascular settings, bulky and stiff wire-type electrodes can intensify sustained abrasion against the compliant endothelial surface, promoting local injury that increases the risk of thrombus formation and vessel narrowing. Even in wearable platforms, prolonged use may cause user discomfort and inflammatory skin symptoms (e.g., dermatitis) 220 , and in severe cases, persistent irritation and local inflammation can precipitate unintended device dislodgement or extrusion 221 . Importantly, even when interface materials are nominally non-cytotoxic and chemically stable, cyclic mechanical loading in harsh in vivo environments can still induce fatigue, degradation, or fracture. Such failure modes may release metal ions and particulate debris, or lead to secondary functional issues such as insulation breakdown and loss of electrode performance, thereby accelerating interface malfunction and indirectly amplifying immune activation 222 ( Fig. 9a , b ). Accordingly, this section highlights four complementary approaches to improve bioelectronic interface biocompatibility: mechanical adjustment, chemical modification, bioactive modification, and drug-releasing surfaces. Fig. 9. Open in a new tab Foreign-body response (FBR) formation of tissue-interface, representative bioelectronic interface contexts, and interfacial engineering strategies. (a) Schematic timeline of a generalized host response at an implanted material/device surface, progressing from Phase I: protein deposition (initial conditioning film and early neutrophil involvement), to Phase II: acute immune activation (monocyte recruitment and macrophage polarization toward M1/M2 phenotypes), Phase III: chronic immune activation (foreign-body giant cell formation and sustained inflammatory milieu), and Phase IV: collagen deposition and fibrosis driven by fibroblasts/myofibroblasts, culminating in a fibrotic capsule that can impair long-term device–tissue coupling. (b) Representative bioelectronic interface contexts covered in this review include neural interfaces, vascular interfaces, epidermal wearable devices, and organ-surface bioelectronics. (c) Overview of interfacial engineering strategies: mechanical adjustment (e.g., miniaturization/compliant structures) to reduce micromotion and tissue injury, chemical modification to regulate protein adsorption and surface-mediated immune activation, bioactive modification to promote cell-selective interaction and tissue integration, and drug-releasing surfaces to locally modulate inflammation and fibrotic remodeling—particularly during early post-implantation phases. 6.1. Mechanical adjustment Conventional electronics built on Si and other inorganic functional materials are typically orders of magnitude stiffer than soft tissues, and they also differ markedly in thickness, deformability, and overall geometry. Such physical mismatches amplify interfacial micromotion and local strain concentrations, which can cause repetitive micro-injury, compromise recording/stimulation fidelity, and exacerbate the foreign-body response (FBR). Immunology studies consistently indicate that large disparities in stiffness and device dimensions are associated with more severe chronic inflammation and encapsulation. Mechanically induced immune responses often originate from interfacial stress caused by mechanical mismatch. Consequently, reducing device rigidity to minimize interfacial stress has become a central design principle, similar to strategies used for achieving conformal interfaces. Broadly, these approaches can be categorized into structural engineering and materials development ( Fig. 9c ). Structure engineering of established electronic materials to achieve effective softness. Instead of changing the functional materials, structural designs—such as macro-/micro-/nano-porosity and mesh-like architectures—reduce the effective flexural rigidity while maintaining electrical performance 107 , 223 . Open, mesh-based implants can minimize tissue trauma and scarring, decrease the total implant “footprint” and flexural rigidity, and facilitate tissue ingrowth and nutrient diffusion 107 , 223 , 224 . Development of intrinsically compliant electronic materials. A complementary route is to replace or hybridize conventional conductors and encapsulants with materials whose mechanical properties better match those of tissue. Hydrogels are particularly attractive because their low modulus and high water content can substantially mitigate stiffness-driven FBR in neural interfaces 225 – 227 . Viscoelastic alginate matrices combined with Ag flakes have enabled conductive hydrogel composites for skin-mounted neuromuscular stimulation electrodes 228 and Ag–polyacrylamide–alginate hydrogels have likewise been used to form conformal interfaces 229 . 6.2. Chemical modification Beyond mechanically driven interfacial trauma, the chemical nature of the electrode and encapsulation materials is also a key determinant of device usability and operational lifetime, because it governs how the interface interacts with the extracellular milieu and immune system. In practice, chemical compatibility is often evaluated along closely related dimensions, including intrinsic cytotoxicity and the propensity to trigger immune-mediated signaling at the material–tissue boundary 222 , 230 , 231 . Classical electronic materials—such as many implant-grade metals and silicon—have been widely adopted in neural electrodes over past decades, in part because they generally exhibit relatively low baseline toxicity in vivo. By contrast, newly developed polymeric systems (especially those that rely on reactive chemistries, initiators, plasticizers, or incomplete curing) can face additional scrutiny regarding material-intrinsic toxicity and the risk of residual monomers or other leachable species. Accordingly, before any implantable interface is advanced to in vivo testing, a comprehensive cytotoxicity and extractables/leachables assessment should be performed to minimize risks associated with material composition and potential unreacted or degradative byproducts 232 , 233 ( Fig. 9c ). In addition to material selection, surface chemical modification of polymer coatings on the device offers an effective route to mitigate FBR 234 . The central idea is to tune the interfacial chemistry such that immune-related signaling cascades are weakened—often by suppressing the earliest events that initiate downstream inflammation and fibrotic remodeling. Because protein adsorption and subsequent cell adhesion are common initiating steps for chronic immune activation, forming a stable, dense, and strongly hydrated antifouling layer at the interface can function as an anti-adhesion barrier. Such a hydration-rich layer reduces nonspecific adsorption of proteins, thereby lowering the likelihood of surface “conditioning” cellular contamination, and the subsequent amplification of immune cell recruitment and proliferation around the implant 236 . For bioelectronic interfaces, an additional constraint is that surface treatments should preserve electrical and ionic coupling between the device and tissue. Therefore, a primary design requirement is to minimize any adverse impact on signal transmission, for which ultrathin antifouling coatings are particularly advantageous. Common antifouling chemistries include poly(ethylene glycol) (PEG) and, more recently, zwitterionic polymer coatings 236 , 237 , both which leverage strong surface hydration to resist protein attachment and can reduce collagen deposition and biofouling for extended periods after implantation. For example, zwitterionic coatings on implanted glucose sensors maintained a low-fouling surface without dense collagenous encapsulation for at least three months, illustrating the potential of charge-balanced, superhydrophilic interfaces to sustain low-fouling conditions in vivo 238 . 6.3. Bioactive modification For certain applications—particularly neural interfaces—bioactive coatings can actively promote long-term compatibility by encouraging neurite ingrowth and supporting neuronal survival near the implant 81 – 83 . Incorporating extracellular-matrix (ECM) components, adhesion ligands, or growth-factor-related cues can modulate the local biological response, while cell-recognition elements (for example, selective antibodies) can enable preferential interactions with specific cell types. These strategies expand the functional scope of bioelectronic interfaces beyond passive toward more biologically integrated, potentially disease-modulating interventions ( Fig. 9c ). ECM-derived bioactive coatings have been used to reduce glial scarring without compromising electrode operation. For instance, laminin coated microelectrode arrays (MEAs) exhibited a pronounced reduction in glial scar formation in a rat model over a one-month implantation period 239 . Similarly, astrocyte-derived ECM coatings of Si MEAs attenuated gliosis relative to clinically used, collagen-based formulations approved by the U.S. Food and Drug Administration (FDA) 83 . Hydrogels are particularly attractive bioactive coatings because their mechanical properties and chemical functionalities are tunable and they can provide a permissive 3D microenvironment that supports cell survival and outgrowth. This versatility enables “living” coatings that incorporate viable cells and exploit their paracrine signaling 226 . For example, hybrid microelectrodes seeded with live hippocampal neurons or astrocytes and subsequently protected by a biodegradable fibrin hydrogel layer substantially reduced reactive astrocyte adsorption and weakened FBR-associated reactions while maintaining the electrode’s electrical resistance 88 . A more targeted direction is to integrate cell-targeting bioactive layers that guide selective interactions with defined cell populations. In one design, SU-8 based mesh electronics were functionalized with antibodies (anti-EAAT2, anti-CD11b, and anti-D2DR) together with a laminin-1 derived synthetic peptide. Following implantation into the mouse hippocampus, chronic electrophysiological measurements and longitudinal histology demonstrated preferential targeting of distinct cell types (neurons, astrocytes, and microglia) and even neuronal subtypes, enabling cell-resolved interfacing that is difficult to achieve with non-functionalized probes 82 . 6.4. Drug-releasing surfaces A complementary strategy to attenuate foreign-body response (FBR) and subsequent scar formation is local immunomodulation, in which immunosuppressive or anti-inflammatory agents are incorporated at or within the bioelectronic–tissue interface. Representative examples include corticosteroids such as dexamethasone and anti-fibrotic compounds. Short-term local administration of these agents can temporarily suppress inflammatory-cell recruitment and, in some contexts, improve protein secretion from immune-isolated cell grafts by reducing immune-mediated isolation. However, many anti-inflammatory drugs are inherently pleiotropic, with broad biological effects and potential dose-limiting toxicities, which can introduce biocompatibility trade-offs beyond mechanics. This makes targeted immunomodulation preferable whenever feasible ( Fig. 9c ). Beyond the four core approaches above, additional interfacial parameters can also shape tissue responses and long-term compatibility, including surface topography (hydrophobic/hydrophilic balance, roughness, and the arrangement/density), surface charge, wettability, and overall implant geometry 240 , 241 . For roughness, submicron features (<1 μm) appear to exert minimal influence on FBR, whereas features on the order of 1–4 μm can improve implant integration in some settings. Spatially confined architectures—where characteristic dimensions are smaller than immune-cell size—may restrict the spreading and activation of pro-inflammatory cells on the material surface 242 , 243 . Importantly, while topographical tuning can be beneficial, it should not compromise the device’s global mechanical integrity or its intended functional performance. Surface charge is another key variable because it can regulate protein adsorption and shape immune-cell–material interactions across different phases of FBR. Adsorption behavior is often governed by net surface charge rather than by purely atomistic electrostatic interactions 244 – 246 . The interfacial charge state is determined by the relationship between the surface isoelectric point and the surrounding fluid pH: pH values below or above the isoelectric point typically yield net positive or negative surface charge, respectively 247 . Electrostatic interactions between negatively charged cells and charged material surfaces can then further modulate immune activation. In general, negatively charged surfaces are frequently associated with a milder response than positively charged surfaces, including thinner encapsulating layers and more limited neovascularization, whereas near-neutral surfaces can reduce protein adsorption and thereby markedly suppress FBR 248 – 250 . In practice, surface charge is commonly evaluated in conjunction with wettability, with protein adsorption serving as an experimentally tractable readout of combined interfacial effects. Finally, because tissue microenvironments vary substantially across implantation sites, biocompatibility should be evaluated in a site-specific manner 251 . 6.5. Comparative framework for selecting biocompatibility strategies Although multiple strategies have been developed to improve the biocompatibility of bioelectronic interfaces, no single approach is universally optimal ( Table 5 ). The most suitable strategy depends on the dominant interfacial challenge in a given application, including mechanical mismatch, surface-mediated immune activation, biofouling, or insufficient biological integration 21 , 252 – 255 . Table 5. Case-based selection of strategies for improving the biocompatibility of tissue–bioelectronics interfaces. Applications Dominant interfacial challenge Preferred strategy Representative evaluation markers Key considerations Neural interfaces Mechanical mismatch, micromotion, and glial scar formation Mechanical adjustment + bioactive modification GFAP/Iba1/CD68 expression, neuronal density, fibrotic/glial scar thickness, impedance stability, signal-to-noise ratio • Ultra-soft materials improve compatibility • Reduce mechanical robustness/Long-term robustness • Improve fabrication/insertion difficulty Vascular interfaces Endothelial injury, Thrombosis, inflammatory response Mechanical compliance + chemical antifouling coatings Platelet adhesion, endothelial integrity, inflammatory infiltration, lumen patency Surface coatings must not compromise sensing or stimulation efficiency Epidermal wearable devices Skin irritation, motion-induced detachment, signal instability Mechanical compliance + interface adhesion engineering Skin erythema score, adhesion stability, motion artifacts, signal reliability Balance between adhesion strength and skin comfort Organ surface bioelectronics Wet adhesion difficulty, fibrosis, immune activation Chemical modification + bioactive interfaces Local inflammation markers, collagen deposition, macrophage infiltration, adhesion strength Wet adhesion and mechanical compliance are both critical Therapeutic implantable bioelectronics Acute inflammatory response after implantation Drug-releasing surfaces + mechanical optimization Cytokine levels, immune-cell infiltration, fibrotic capsule thickness, device functional stability Drug release often effective for shortterm modulation Open in a new tab Mechanical adjustment is particularly critical for chronically implanted devices interfacing with soft and dynamic tissues, where stiffness mismatch and micromotion can drive tissue injury and foreign-body responses. Chemical surface modification is especially beneficial for interfaces limited by protein adsorption and biofouling, such as electrochemical sensors or long-term fluid-exposed devices. Bioactive modification strategies are advantageous when active tissue integration or cell-selective interaction is required, for example in neural or regenerative biointerfaces. Drug-releasing surfaces can provide local immunomodulation during the acute post-implantation phase but are often limited by finite release duration and potential off-target effects 253 , 255 . In practice, these approaches are often most effective when combined, where structural engineering, chemical interface regulation, and bioactive modulation collectively shape the tissue–device interface 21 , 252 – 254 . In addition, the effectiveness of biocompatibility optimization strategies needs to be evaluated using a combination of material-level, biological, and device-level markers. At the material and interface level, relevant parameters include mechanical properties (e.g., modulus and bending stiffness), surface wettability, protein adsorption, coating stability, and electrical interfacial properties such as impedance and charge transfer efficiency 21 , 252 , 254 , 256 . Biological responses can be assessed through acute and chronic markers of host response, including cell viability, inflammatory cytokines (e.g., TNF-α, IL-1β, and IL-6), immune-cell infiltration, macrophage activation, and fibrotic capsule formation 253 , 255 , 256 . For neural interfaces, markers of glial activation and neuronal preservation are also commonly used. Finally, device-level functional metrics provide important evidence of improved biocompatibility. These include signal-to-noise ratio, impedance stability over time, stimulation threshold, sensing stability, and long-term recording fidelity 252 . Together, these markers enable a comprehensive evaluation of how different strategies influence both biological responses and device performance. 7. Outlook Future progress in tissue–bioelectronics interfaces will rely on deeper integration of materials design, structural engineering and biological insight. Although substantial progress has enabled soft, conformal, and multifunctional interfaces, achieving durable and predictable performance within the dynamic, heterogeneous tissue microenvironment remains a central challenge. Several emerging directions are expected to address key limitations in long-term stability, interfacial adaptation, and system integration. Adaptive materials capable of tuning mechanical and chemical properties in vivo may help accommodate tissue remodeling over time; biohybrid and immunomodulatory interfaces that actively engage with host tissues offer pathways toward long-term functional integration rather than passive tolerance. In parallel, computational and data-driven design frameworks may enable predictive optimization of interfacial mechanics, adhesion energetics, and immune responses. Advances in 3D and in vivo printing, personalized device architectures, and minimally invasive deployment strategies will further expand the design space for interfaces tailored to complex anatomical geometries. Collectively, these developments indicate a transition toward bioelectronic systems that dynamically adapt to tissue remodeling, maintain stable coupling, and enable feedback-controlled operation. Realizing this vision, however, will require more than improvements in short-term biocompatibility or proof-of-concept functionality. Clinical translation also depends on whether device fate, interface durability, system reliability, manufacturability, and regulatory readiness can be addressed across the full lifecycle of tissue-interfaced bioelectronics. 7.1. Clinical translation: device fate and end-of-life considerations Beyond short-term tissue compatibility and proof-of-concept functionality, clinical translation requires explicit consideration of the long-term fate and end-of-life scenarios of tissue-interfaced bioelectronics ( Table 6 ). For long-term or permanent implants, device fate is often governed by chronic interfacial evolution (e.g., fibrotic encapsulation and biofouling) and mechanical/electrochemical failure modes such as fatigue, delamination, insulation breakdown, and corrosion, which can progressively compromise both safety and function. In mechanically harsh in vivo environments, failure may also lead to the release of metal ions or particulate debris, indirectly amplifying immune activation and accelerating functional decay 21 , 256 , 257 . Table 6. Translational implications of conformal-interface design levers in structural engineering and materials viscoelasticity. Structure engineering and material design principles Intended interfacial mechanism Typical use scenario Device fate & end-of-life implications Translation-focused validation tests Regulatory / manufacturing watch-outs Structural engineering: ultrathin/low bending stiffness (ultrathin films, thin encapsulation) Reduce effective stiffness; suppress micromotion-induced injury; improve conformal contact Chronic soft-tissue interfaces (neural/organ-surface), longterm wearable sensing Long-term operation possible, but handling/implantation damage risk; thin barriers can accelerate moisture ingress and functional drift Wet-state mechanical durability (bend/strain cycling); encapsulation integrity in physiological media; long-term impedance drift/SNR stability; accelerated aging Process window tight; thickness uniformity & pinhole defects; packaging/sterilization may change barrier/adhesion Structural engineering: strain-accommodating interconnects (serpentine/island-bridge) Decouple rigid components from tissue strain; mitigate lead fracture Hybrid systems: rigid chips + soft substrate; stretchable arrays End-of-life often dominated by interconnect fatigue or connector failure; revision may require modular explant Cyclic fatigue under multiaxial deformation; electrical continuity under motion; connector pull/peel reliability; failure-mode analysis post-test QC for microcracks/line breaks; batch-to-batch geometry tolerance; assembly yield and traceability Structural engineering: out-of-plane / 3D compliant architectures (mesh, porous, textile-like) Promote tissue compliance and reduce stress concentration; (sometimes) enable tissue integration Neural probes/soft implants; long-term interfaces requiring stable placement May lead to tissue ingrowth → improved stability but harder retrieval/revision; end-of-life planning becomes essential Chronic implantation stability; explant feasibility; histology + functional stability over time; imaging compatibility Classification and surgical workflow complexity; manufacturing consistency of 3D geometry; sterilization penetration and packaging robustness Structural engineering: geometrical cuts / kirigami-origami Allow large stretchability without high material strain; reduce interface shear Highly deformable wearables; dynamic organ surfaces Fate often limited by tear initiation at cut tips; mechanical damage can trigger delamination and signal drift Tear/fatigue resistance at stress concentrators; adhesion retention under repeated stretch; motion-artifact quantification Defect sensitivity high; inspection standards needed; reproducibility of cut features at scale Elastic materials (stable low-modulus elastomers) Provide immediate compliance with predictable elastic response; maintain geometry over time Long-term wearables; encapsulation layers; soft substrates Generally compatible with long-term residence, but aging can cause stiffening/softening, changing interface stress & sealing Long-term mechanical aging (creep/relaxation); encapsulation/barrier aging; impedance stability; sterilization compatibility Formulation control; sterilization-induced property drift; shelf-life and packaging constraints Viscoelastic materials (stress-relaxing hydrogels/ionogels) Time-dependent stress relaxation to reduce chronic irritation; improve wet contact Soft tissue/organ interfaces; hydrogel-based electrodes Fate depends on swelling/dehydration, leaching, and network remodeling; risk of progressive property drift affecting function Swelling/dehydration cycling; extractables/leachables; long-term adhesion in wet; electrical drift with hydration changes Batch control of crosslink density & water content; storage/transport humidity; sterilization effects on network chemistry Viscous / flowable materials (pastes/flowable inks/in situ forming layers) Immediate intimate contact; fill microgaps; conform to rough surfaces Acute interfaces; temporary fixation; in situ formed/printed contact layers Fate includes migration/redistribution, incomplete curing, residual species; retrievability/revision becomes key Curing completeness; residual/reactive species + leachables; in situ geometry reproducibility; migration/dispersion assessment Operator dependence; workflow validation; stronger regulatory scrutiny for in situ formation; QC for curing and composition Cross-cutting: hybrid + multifunctional systems (sensor/stimulator + coatings/drug layers) Combine mechanical conformity + biochemical modulation + functionality Closed-loop therapies; drug-eluting or bioactive interfaces Fate may be governed by the “weakest link”: coating depletion, drug exhaustion, mechanical fatigue, or electrical drift Integrated safety + performance endpoints (biology + function); chronic stability; post-explant failure analysis Possible combination-product classification; manufacturing traceability; consistency of multi-layer stacks; sterilization/shelf-life across components Open in a new tab For transient or resorbable bioelectronics, additional considerations arise regarding degradation kinetics and mechanisms (surface vs bulk erosion), local accumulation and systemic clearance of degradation byproducts, and potential mismatches between functional lifetime and physical resorption time. These issues become particularly important for devices that incorporate conductive fillers, nanoparticles, or hybrid material stacks, where partial degradation may leave residual fragments that persist longer than intended. For in situ formed/printed interfaces, translational feasibility further depends on curing completeness, geometric reproducibility, and whether the device can be retrieved or revised in case of malfunction 258 – 260 . 7.2. Translational testing: from materials to interface durability and system reliability A clinically relevant evaluation pathway should extend beyond standard cytotoxicity screening to include staged validation across material-, interface-, and system-level performance. At the material level, early testing typically includes cytotoxicity as well as extractables/leachables assessments to de-risk residual monomers, initiators, plasticizers, and degradation byproducts, especially for emerging polymeric and reactive chemistries. At the interface level, translation-oriented testing should emphasize durability under wet and dynamic physiological conditions, including adhesion retention, delamination resistance, and mechanical fatigue under cyclic deformation. For bioelectronic interfaces, electrical stability metrics (e.g., impedance drift, signal-to-noise ratio stability, stimulation threshold drift, baseline drift for electrochemical sensors) provide direct evidence of whether improved biocompatibility translates into sustained functional coupling 21 , 261 , 262 . Practical constraints such as sterilization compatibility, packaging, and storage/shelf-life stability are also critical, as these processes can alter surface chemistries, hydration layers, or reactive functional groups, thereby changing interfacial behavior prior to implantation 257 . In vivo validation ideally progresses from acute to chronic small-animal studies and, when feasible, to larger-animal models that better capture anatomical scale, surgical workflow, device fixation, motion environment, and revision feasibility. Post-explant analysis, including failure-mode characterization and tissue-level histology, can further inform design iteration by linking interfacial biological response to device degradation, mechanical damage, and electrical drift 257 , 260 . 7.3. Regulatory and manufacturing challenges for clinically deployable bioelectronic interfaces Regulatory translation is complicated by the rapidly increasing material and functional complexity of tissue–bioelectronics interfaces. Many emerging systems incorporate unconventional material chemistries (reactive adhesives, bioactive coatings, degradable conductors), hybrid integration of rigid and soft components, and in some cases drug-releasing or closed-loop therapeutic functions. These features can complicate device classification and increase the burden of evidence required to demonstrate safety and effectiveness, particularly when long-term exposure, degradation byproducts, or residual reactive species are involved 260 , 263 , 264 . Equally important, manufacturing reproducibility and quality control represent regulatory-relevant challenges rather than purely engineering concerns. For interfaces relying on micro/nano-structured designs or time-dependent material behaviors, batch-to-batch variability in geometry, curing extent, hydration state, or filler dispersion can lead to substantial variance in both biological response and electrical performance. Establishing standardized evaluation protocols that integrate biological safety, interfacial stability, electrical reliability, sterilization/shelf-life constraints, and end-of-life considerations will be critical for advancing tissue–bioelectronics interfaces from laboratory prototypes to clinically robust technologies 257 , 260 – 262 . Table 3. Representative therapeutic bioelectronic systems. Name Device format Interfacial strategy Material Fabrication techniques Interface implication Practical constraint / coupling condition Wireless closed-loop smart bandage 206 Wearable therapy Mesh + island–bridge (+ microstructure) (strain-tolerant stretchable electronics + patterned sensing/actuation; wet-interface design) Hydrogel + rubber (hybrid) Stretchable microfabrication/transfer of serpentine circuits → multilayer sealing → hydrogel electrode/interface integration + wireless packaging Conformal wet adhesion and strain-tolerant routing enable stable sensing–actuation coupling in exudate-rich wound environments. Biofouling/sealing integrity and long-term hydrogel/adhesive stability largely determine drift and reliability during repeated use. Electronic dura mater 268 Implantable (spinal/brain surface) Ultrathin + microstructure (+ mesh) (conformal silicone implant with stretchable interconnects + microfluidics) Rubber (silicone) Soft-lithography microfluidic layers + thin-film patterned stretchable interconnects/electrodes → multilayer bonding → connect Dura-matched compliance and integrated delivery/stimulation channels reduce micromotion injury while enabling multimodal neuromodulation. Multi-material integration and terminal connect/packaging govern chronic wet stability and functional longevity. Gastric-retentive robotic capsule 269 Ingestible therapy Microstructure (retention + controlled release actuation modules) Other + (liquid payload) (rigid capsule + liquid reservoir) Precision capsule assembly integrating retention/actuation modules → sealed reservoir/exposure control → telemetry/power integration Retention and actuation modules define exposure control and enable responsive dosing without a permanent implant footprint. Safety of long residence, power/telemetry robustness, and retention control dominate practical deployment and regulation. Bioabsorbable mechanoelectrical stimulation suture 270 Implantable (suture-like) Microstructure (procedure-compatible thread interface; selective exposure at wound line) Rubber-like solid + paste (hybrid) Coaxial fiber assembly/layering → define exposed stimulation region → lifetime-tuned biodegradable encapsulation/packaging Thread-like geometry enables therapy to be delivered at the wound line with minimal added bulk and workflow compatibility. Output depends on local strain/motion and on predictable wet degradation; terminal exposure/encapsulation must remain stable. Body-coupled stimulation socks 271 Wearable therapy Mesh (garment-scale conductive routing + distributed electrodes) Rubber-like solid (textile network) Textile integration of conductive routing/electrodes → localized reinforcement/encapsulation of traces → module connection Garment-scale routing distributes stimulation over large areas while preserving comfort and breathability. Contact pressure variability and wash/sweat durability (routing + connectors + selective protection) largely determine reproducibility. Open in a new tab Acknowledgements This work was supported through National Institutes of Health grants R01HL155815 and R01DC021461, National Science Foundation grants 2145802 and 2444815, US Army Medical Research Acquisition Activity grant HT9425–24-1–0249, Army Research Office grant W911NF-23–1-0041, Office of Naval Research grant N00014–25-1–2258, and Heritage Medical Research Institute. Footnotes Competing interests The authors declare no competing interests. Data availability No primary research data were generated or analyzed as part of this Review. All information discussed is derived from previously published studies, which are cited within the manuscript. References 1. Chen C, Ding S & Wang J Digital health for aging populations. Nat. Med. 29, 1623–1630 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 2. Hajat C & Stein E The global burden of multiple chronic conditions: A narrative review. Prev. Med. Rep. 12, 284–293 (2018). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 3. Madhvapathy SR et al. Implantable bioelectronics and wearable sensors for kidney health and disease. Nat. Rev. Nephrol. 21, 443–463 (2025). [ DOI ] [ PubMed ] [ Google Scholar ] 4. Brasier N et al. Applied body-fluid analysis by wearable devices. Nature 636, 57–68 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 5. Min S et al. Wearable blood pressure sensors for cardiovascular monitoring and machine learning algorithms for blood pressure estimation. Nat. Rev. Cardiol. 22, 629–648 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 6. Wang C et al. Wound management materials and technologies from bench to bedside and beyond. Nat. Rev. Mater. 9, 550–566 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 7. Perkins BA, Sherr JL & Mathieu C Type 1 diabetes glycemic management: Insulin therapy, glucose monitoring, and automation. Science 373, 522–527 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 8. Chen R, Canales A & Anikeeva P Neural recording and modulation technologies. Nat. Rev. Mater. 2, 16093 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 9. Barrese JC et al. Failure mode analysis of silicon-based intracortical microelectrode arrays in non-human primates. J. Neural Eng. 10, 066014 (2013). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 10. Wynn TA & Ramalingam TR Mechanisms of fibrosis: therapeutic translation for fibrotic disease. Nat. Med. 18, 1028–1040 (2012). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 11. Tang X, Shen H, Zhao S, Li N & Liu J Flexible brain–computer interfaces. Nat. Electron. 6, 109–118 (2023). [ Google Scholar ] 12. Ni Y et al. Macrophages modulate stiffness-related foreign body responses through plasma membrane deformation. Proc. Natl. Acad. Sci. 120, e2213837120 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 13. Martino F, Perestrelo AR, Vinarský V, Pagliari S & Forte G Cellular Mechanotransduction: From Tension to Function. Front. Physiol. 9, (2018). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 14. Jaalouk DE & Lammerding J Mechanotransduction gone awry. Nat. Rev. Mol. Cell Biol. 10, 63–73 (2009). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 15. Anderson JM, Rodriguez A & Chang DT Foreign body reaction to biomaterials. Semin. Immunol. 20, 86–100 (2008). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 16. Tang L & Eaton JW Fibrin(ogen) mediates acute inflammatory responses to biomaterials. J. Exp. Med. 178, 2147–2156 (1993). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 17. Fujiwara N & Kobayashi K Macrophages in Inflammation. Curr. Drug Targets Inflamm. Allergy 4, 281–286 (2005). [ DOI ] [ PubMed ] [ Google Scholar ] 18. Cogan SF Neural Stimulation and Recording Electrodes. Annu. Rev. Biomed. Eng. 10, 275–309 (2008). [ DOI ] [ PubMed ] [ Google Scholar ] 19. Madhvapathy SR et al. Implantable bioelectronic systems for early detection of kidney transplant rejection. Science 381, 1105–1112 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 20. Kim D-H et al. Dissolvable films of silk fibroin for ultrathin conformal bio-integrated electronics. Nat. Mater. 9, 511–517 (2010). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 21. Song E, Li J, Won SM, Bai W & Rogers JA Materials for flexible bioelectronic systems as chronic neural interfaces. Nat. Mater. 19, 590–603 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 22. Rogers JA, Someya T & Huang Y Materials and Mechanics for Stretchable Electronics. Science 327, 1603–1607 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 23. Nishio Y et al. Intrinsically stretchable transistors and integrated circuits. Nat. Rev. Electr. Eng. 2, 715–735 (2025). [ Google Scholar ] 24. Zhao X et al. Soft Materials by Design: Unconventional Polymer Networks Give Extreme Properties. Chem. Rev. 121, 4309–4372 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 25. Vegas AJ et al. Combinatorial hydrogel library enables identification of materials that mitigate the foreign body response in primates. Nat. Biotechnol. 34, 345–352 (2016). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 26. Kim D-H et al. Epidermal Electronics. Science 333, 838–843 (2011). [ DOI ] [ PubMed ] [ Google Scholar ] 27. Nam S & Mooney D Polymeric Tissue Adhesives. Chem. Rev. 121, 11336–11384 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 28. Boufidis D, Garg R, Angelopoulos E, Cullen DK & Vitale F Bio-inspired electronics: Soft, biohybrid, and “living” neural interfaces. Nat. Commun. 16, 1861 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 29. Wang X et al. Tough Wet Adhesion of Hydrogen-Bond-Based Hydrogel with On-Demand Debonding and Efficient Hemostasis. ACS Appl. Mater. Interfaces 14, 36166–36177 (2022). [ DOI ] [ PubMed ] [ Google Scholar ] 30. Kurokawa T, Furukawa H, Wang W, Tanaka Y & Gong JP Formation of a strong hydrogel–porous solid interface via the double-network principle. Acta Biomater. 6, 1353–1359 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 31. Reilly A & Tkatchenko A van der Waals dispersion interactions in molecular materials: beyond pairwise additivity. Chem. Sci. 6, 3289–3301 (2015). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 32. Yan Z et al. Highly stretchable van der Waals thin films for adaptable and breathable electronic membranes. Science 375, 852–859 (2022). [ DOI ] [ PubMed ] [ Google Scholar ] 33. Nie B, Stutzman J & Xie A A Vibrational Spectral Maker for Probing the Hydrogen-Bonding Status of Protonated Asp and Glu Residues. Biophys. J. 88, 2833–2847 (2005). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 34. Chen J et al. An Adhesive Hydrogel with “Load-Sharing” Effect as Tissue Bandages for Drug and Cell Delivery. Adv. Mater. 32, 2001628 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 35. Yuk H et al. Dry double-sided tape for adhesion of wet tissues and devices. Nature 575, 169–174 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 36. Xue B et al. Hydrogel tapes for fault-tolerant strong wet adhesion. Nat. Commun. 12, 7156 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 37. Kim J et al. Soft Cardiac Patch Using a Bifacial Architecture of Adhesive/Low-Impedance Hydrogel Nanocomposites and Highly Conductive Elastomer Nanocomposites. Adv. NanoBiomed Res. 5, 2400143 (2025). [ Google Scholar ] 38. Palacio MLB & Bhushan B Bioadhesion: a review of concepts and applications. Philos. Trans. R. Soc. Math. Phys. Eng. Sci. 370, 2321–2347 (2012). [ DOI ] [ PubMed ] [ Google Scholar ] 39. Rao P et al. Tough hydrogels with fast, strong, and reversible underwater adhesion based on a multiscale design. Adv. Mater. 30, 1801884 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 40. Roy CK et al. Self-Adjustable Adhesion of Polyampholyte Hydrogels. Adv. Mater. 27, 7344–7348 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 41. Montazerian H et al. Stretchable and Bioadhesive Gelatin Methacryloyl-Based Hydrogels Enabled by in Situ Dopamine Polymerization. ACS Appl. Mater. Interfaces 13, 40290–40301 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 42. Li J et al. Tough adhesives for diverse wet surfaces. Science 357, 378–381 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 43. Yang Q et al. Photocurable bioresorbable adhesives as functional interfaces between flexible bioelectronic devices and soft biological tissues. Nat. Mater. 20, 1559–1570 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 44. Young T III. An essay on the cohesion of fluids. Philos. Trans. R. Soc. Lond. 65–87 (1805) doi: 10.1098/rstl.1805.0005. [ DOI ] [ Google Scholar ] 45. Delamarche E et al. Microfluidic Networks for Chemical Patterning of Substrates:  Design and Application to Bioassays. J. Am. Chem. Soc. 120, 500–508 (1998). [ Google Scholar ] 46. Wisniewski DJ et al. Spatial control of doping in conducting polymers enables complementary, conformable, implantable internal ion-gated organic electrochemical transistors. Nat. Commun. 16, 517 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 47. Boys AJ et al. Implantable bioelectronics for gut electrophysiology. Nat. Commun. 16, 10240 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 48. Xu D et al. Two-dimensional semiconductor-based active array for high-fidelity spatiotemporal monitoring of neural activities. Nat. Mater. 1–12 (2025) doi: 10.1038/s41563-025-02430-4. [ DOI ] [ PubMed ] [ Google Scholar ] 49. Park H-H et al. Flexible and Shape-Reconfigurable Hydrogel Interlocking Adhesives for High Adhesion in Wet Environments Based on Anisotropic Swelling of Hydrogel Microstructures. ACS Macro Lett. 6, 1325–1330 (2017). [ DOI ] [ PubMed ] [ Google Scholar ] 50. Wang Y, Kang V, Arzt E, Federle W & Hensel R Strong Wet and Dry Adhesion by Cupped Microstructures. ACS Appl. Mater. Interfaces 11, 26483–26490 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 51. Yang J, Bai R, Chen B & Suo Z Hydrogel Adhesion: A Supramolecular Synergy of Chemistry, Topology, and Mechanics. Adv. Funct. Mater. 30, 1901693 (2020). [ Google Scholar ] 52. Samieipour F, Dianat-Moghadam H & Khanahmad H Recent developments in bioconjugation: From strategies to design and clinical applications. Biomed. Pharmacother. 192, 118593 (2025). [ DOI ] [ PubMed ] [ Google Scholar ] 53. Yuk H, Zhang T, Lin S, Parada GA & Zhao X Tough bonding of hydrogels to diverse non-porous surfaces. Nat. Mater. 15, 190–196 (2016). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 54. Wallace DG et al. A tissue sealant based on reactive multifunctional polyethylene glycol. J. Biomed. Mater. Res. 58, 545–555 (2001). [ DOI ] [ PubMed ] [ Google Scholar ] 55. Preul MC, Campbell PK, Bichard WD & Spetzler RF Application of a hydrogel sealant improves watertight closures of duraplasty onlay grafts in a canine craniotomy model. 107, (2007). [ DOI ] [ PubMed ] [ Google Scholar ] 56. Artzi N, Shazly T, Baker AB, Bon A & Edelman ER Aldehyde-Amine Chemistry Enables Modulated Biosealants with Tissue-Specific Adhesion. Adv. Mater. Deerfield Beach Fla 21, 3399–3403 (2009). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 57. Lin H et al. Synthesis and characterization of a bio-aldehyde-based lignin adhesive with desirable water resistance. Int. J. Biol. Macromol. 264, 130020 (2024). [ DOI ] [ PubMed ] [ Google Scholar ] 58. Chao H-H & Torchiana DF BioGlue: Albumin/Glutaraldehyde Sealant in Cardiac Surgery. J. Card. Surg. 18, 500–503 (2003). [ DOI ] [ PubMed ] [ Google Scholar ] 59. Belowich ME & Stoddart JF Dynamic imine chemistry. Chem. Soc. Rev. 41, 2003–2024 (2012). [ DOI ] [ PubMed ] [ Google Scholar ] 60. M. Korde J & Kandasubramanian B. Biocompatible alkyl cyanoacrylates and their derivatives as bio-adhesives. Biomater. Sci. 6, 1691–1711 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 61. Klemarczyk P The isolation of a zwitterionic initiating species for ethyl cyanoacrylate (ECA) polymerization and the identification of the reaction products between 1°, 2°, and 3° amines with ECA. Polymer 42, 2837–2848 (2001). [ Google Scholar ] 62. Garcia-Roig M, Gorin MA, Castellan M & Ciancio G OMNEX Surgical Sealant in the Extracorporeal Repair of Renal Artery Aneurysms. Ann. Vasc. Surg. 25, 1141.e5–1141.e8 (2011). [ DOI ] [ PubMed ] [ Google Scholar ] 63. Singer AJ, Perry LC & Allen RL Jr In Vivo Study of Wound Bursting Strength and Compliance of Topical Skin Adhesives. Acad. Emerg. Med. 15, 1290–1294 (2008). [ DOI ] [ PubMed ] [ Google Scholar ] 64. Investigators, T. n-B. T. n-Butyl Cyanoacrylate Embolization of Cerebral Arteriovenous Malformations: Results of a Prospective, Randomized, Multi-center Trial. Am. J. Neuroradiol. 23, 748–755 (2002). [ PMC free article ] [ PubMed ] [ Google Scholar ] 65. Nair LS & Laurencin CT Biodegradable polymers as biomaterials. Prog. Polym. Sci. 32, 762–798 (2007). [ Google Scholar ] 66. Hofman AH, van Hees IA, Yang J & Kamperman M Bioinspired Underwater Adhesives by Using the Supramolecular Toolbox. Adv. Mater. 30, 1704640 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 67. Davoodi E et al. Imaging-guided deep tissue in vivo sound printing. Science 388, 616–623 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 68. Yang J, Stuart MAC & Kamperman M Jack of all trades: versatile catechol crosslinking mechanisms. 43, 8271–8298 (2014). [ DOI ] [ PubMed ] [ Google Scholar ] 69. Lee HA, Park E & Lee H Polydopamine and Its Derivative Surface Chemistry in Material Science: A Focused Review for Studies at KAIST. Adv. Mater. 32, 1907505 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 70. Hoang Thi TT, Lee Y, Le Thi P & Park KD Engineered horseradish peroxidase-catalyzed hydrogels with high tissue adhesiveness for biomedical applications. J. Ind. Eng. Chem. 78, 34–52 (2019). [ Google Scholar ] 71. Ono K et al. Photocrosslinkable chitosan as a biological adhesive. J. Biomed. Mater. Res. 49, 289–295 (2000). [ DOI ] [ PubMed ] [ Google Scholar ] 72. Gilbert TW et al. Lysine-Derived Urethane Surgical Adhesive Prevents Seroma Formation in a Canine Abdominoplasty Model. Plast. Reconstr. Surg. 122, 95 (2008). [ DOI ] [ PubMed ] [ Google Scholar ] 73. Hong G & Lieber CM Novel electrode technologies for neural recordings. Nat. Rev. Neurosci. 20, 330–345 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 74. Lecomte A, Descamps E & Bergaud C A review on mechanical considerations for chronically-implanted neural probes. J. Neural Eng. 15, 031001 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 75. Fu T-M et al. Stable long-term chronic brain mapping at the single-neuron level. Nat. Methods 13, 875–882 (2016). [ DOI ] [ PubMed ] [ Google Scholar ] 76. Luan L et al. Ultraflexible nanoelectronic probes form reliable, glial scar–free neural integration. Sci. Adv. 3, e1601966 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 77. Musk E & Neuralink. An Integrated Brain-Machine Interface Platform With Thousands of Channels. J. Med. Internet Res. 21, e16194 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 78. Zhou T et al. 3D printable high-performance conducting polymer hydrogel for all-hydrogel bioelectronic interfaces. Nat. Mater. 22, 895–902 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 79. Dai Y et al. Soft hydrogel semiconductors with augmented biointeractive functions. Science 386, 431–439 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 80. Dobashi Y et al. Piezoionic mechanoreceptors: Force-induced current generation in hydrogels. Science 376, 502–507 (2022). [ DOI ] [ PubMed ] [ Google Scholar ] 81. Zhu B et al. Large enhancement in neurite outgrowth on a cell membrane-mimicking conducting polymer. Nat. Commun. 5, 4523 (2014). [ DOI ] [ PubMed ] [ Google Scholar ] 82. Zhang A, Zwang TJ & Lieber CM Biochemically functionalized probes for cell-type–specific targeting and recording in the brain. Sci. Adv. 9, eadk1050 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 83. Oakes RS, Polei MD, Skousen JL & Tresco PA An astrocyte derived extracellular matrix coating reduces astrogliosis surrounding chronically implanted microelectrode arrays in rat cortex. Biomaterials 154, 1–11 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 84. He W & Bellamkonda RV Nanoscale neuro-integrative coatings for neural implants. Biomaterials 26, 2983–2990 (2005). [ DOI ] [ PubMed ] [ Google Scholar ] 85. Gearing M & Kennedy P Histological Confirmation of Myelinated Neural Filaments Within the Tip of the Neurotrophic Electrode After a Decade of Neural Recordings. Front. Hum. Neurosci. 14, (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 86. Kennedy PR The cone electrode: a long-term electrode that records from neurites grown onto its recording surface. J. Neurosci. Methods 29, 181–193 (1989). [ DOI ] [ PubMed ] [ Google Scholar ] 87. Rochford AE et al. Functional neurological restoration of amputated peripheral nerve using biohybrid regenerative bioelectronics. Sci. Adv. 9, eadd8162 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 88. De Faveri S et al. Bio-inspired hybrid microelectrodes: a hybrid solution to improve longterm performance of chronic intracortical implants. Front. Neuroengineering 7, (2014). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 89. Purcell EK, Seymour JP, Yandamuri S & Kipke DR In vivo evaluation of a neural stem cell-seeded prosthesis. J. Neural Eng. 6, 026005 (2009). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 90. Miles AW & Gheduzzi S Basic biomechanics and biomaterials. Surg. Oxf. 27, 90–95 (2009). [ Google Scholar ] 91. Widlund T, Yang S, Hsu Y-Y & Lu N Stretchability and compliance of freestanding serpentine-shaped ribbons. Int. J. Solids Struct. 51, 4026–4037 (2014). [ Google Scholar ] 92. Zhang Y et al. Mechanics of ultra-stretchable self-similar serpentine interconnects. Acta Mater. 61, 7816–7827 (2013). [ Google Scholar ] 93. Timoshenko SP & Gere JM Theory of Elastic Stability. (Courier Corporation, 2012). [ Google Scholar ] 94. Fan Z et al. A finite deformation model of planar serpentine interconnects for stretchable electronics. Int. J. Solids Struct. 91, 46–54 (2016). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 95. Dagdeviren C et al. Conformal piezoelectric systems for clinical and experimental characterization of soft tissue biomechanics. Nat. Mater. 14, 728–736 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 96. Kim D-H et al. Stretchable and Foldable Silicon Integrated Circuits. Science 320, 507–511 (2008). [ DOI ] [ PubMed ] [ Google Scholar ] 97. Yang D et al. Double-Microcrack Coupling Stretchable Neural Electrode for Electrophysiological Communication. Adv. Funct. Mater. 33, 2300412 (2023). [ Google Scholar ] 98. Nur R et al. A Highly Sensitive Capacitive-type Strain Sensor Using Wrinkled Ultrathin Gold Films. Nano Lett. 18, 5610–5617 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 99. Song J et al. An analytical study of two-dimensional buckling of thin films on compliant substrates. J. Appl. Phys. 103, 014303 (2008). [ Google Scholar ] 100. Babaee S et al. Kirigami-inspired stents for sustained local delivery of therapeutics. Nat. Mater. 20, 1085–1092 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 101. Pradhan S et al. Biofunctional Silk Kirigami With Engineered Properties. ACS Appl. Mater. Interfaces 12, 12436–12444 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 102. Yu Y et al. All-printed soft human-machine interface for robotic physicochemical sensing. Sci. Robot. 7, eabn0495 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 103. Ko HC et al. A hemispherical electronic eye camera based on compressible silicon optoelectronics. Nature 454, 748–753 (2008). [ DOI ] [ PubMed ] [ Google Scholar ] 104. Kim D-H et al. Materials and noncoplanar mesh designs for integrated circuits with linear elastic responses to extreme mechanical deformations. Proc. Natl. Acad. Sci. 105, 18675–18680 (2008). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 105. Han M et al. Catheter-integrated soft multilayer electronic arrays for multiplexed sensing and actuation during cardiac surgery. Nat. Biomed. Eng. 4, 997–1009 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 106. Someya T et al. Conformable, flexible, large-area networks of pressure and thermal sensors with organic transistor active matrixes. Proc. Natl. Acad. Sci. 102, 12321–12325 (2005). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 107. Zhou T et al. Syringe-injectable mesh electronics integrate seamlessly with minimal chronic immune response in the brain. Proc. Natl. Acad. Sci. 114, 5894–5899 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 108. Zhang A et al. Ultraflexible endovascular probes for brain recording through micrometer-scale vasculature. Science 381, 306–312 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 109. Sun H et al. Liquid-based encapsulation for implantable bioelectronics across broad pH environments. Nat. Commun. 16, 1019 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 110. Ciatti JL et al. An autonomous implantable device for the prevention of death from opioid overdose. Sci. Adv. (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 111. Xu Y et al. Phase-separated porous nanocomposite with ultralow percolation threshold for wireless bioelectronics. Nat. Nanotechnol. 19, 1158–1167 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 112. Deng J et al. Electrical bioadhesive interface for bioelectronics. Nat. Mater. 20, 229–236 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 113. Wang C et al. Bioadhesive ultrasound for long-term continuous imaging of diverse organs. Science 377, 517–523 (2022). [ DOI ] [ PubMed ] [ Google Scholar ] 114. Landau LD, Pitaevskii LP, Kosevich AM & Lifshitz EM Theory of Elasticity: Volume 7. (Elsevier, 2012). [ Google Scholar ] 115. Chaudhuri O et al. Extracellular matrix stiffness and composition jointly regulate the induction of malignant phenotypes in mammary epithelium. Nat. Mater. 13, 970–978 (2014). [ DOI ] [ PubMed ] [ Google Scholar ] 116. Ferry JD & Myers HS Viscoelastic Properties of Polymers. J. Electrochem. Soc. 108, 142C (1961). [ Google Scholar ] 117. Kundu PK, Cohen IM, Dowling DR & Capecelatro J Fluid Mechanics. (Elsevier, 2024). [ Google Scholar ] 118. Kim J et al. A wireless, implantable bioelectronic system for monitoring urinary bladder function following surgical recovery. Proc. Natl. Acad. Sci. 121, e2400868121 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 119. Wang S et al. Skin electronics from scalable fabrication of an intrinsically stretchable transistor array. Nature 555, 83–88 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 120. Sheng H et al. Brain implantation of soft bioelectronics via embryonic development. Nature 642, 954–964 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 121. Choi S et al. Highly conductive, stretchable and biocompatible Ag–Au core–sheath nanowire composite for wearable and implantable bioelectronics. Nat. Nanotechnol. 13, 1048–1056 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 122. Markvicka EJ, Bartlett MD, Huang X & Majidi C An autonomously electrically self-healing liquid metal–elastomer composite for robust soft-matter robotics and electronics. Nat. Mater. 17, 618–624 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 123. Yang J, Cheng W & Kalantar-zadeh K Electronic Skins Based on Liquid Metals. Proc. IEEE 107, 2168–2184 (2019). [ Google Scholar ] 124. Cole T, Khoshmanesh K & Tang S-Y Liquid Metal Enabled Biodevices. Adv. Intell. Syst. 3, 2000275 (2021). [ Google Scholar ] 125. Afzal S et al. Ultrasound Hydrogel: A Review on Materials and Method. J. Mod. Polym. Chem. Mater. 1(1), (2022). [ Google Scholar ] 126. Zhang R, Wen S, Zhao Y & Ji S Water-Removable Paste Electrode for Electrophysiological Monitoring on Hairy Skin. ACS Appl. Electron. Mater. 7, 925–932 (2025). [ Google Scholar ] 127. Wang X et al. Towards Optimizing the Quality of Long-Term Physiological Signals Monitoring by Using Anhydrous Carbon Paste Electrode. IEEE Trans. Biomed. Eng. 70, 423–435 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 128. Tringides CM et al. Viscoelastic surface electrode arrays to interface with viscoelastic tissues. Nat. Nanotechnol. 16, 1019–1029 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 129. Kim HJ, Koo JH, Lee S, Hyeon T & Kim D-H Materials design and integration strategies for soft bioelectronics in digital healthcare. Nat. Rev. Mater. 10, 654–673 (2025). [ Google Scholar ] 130. Dai Y, Hu H, Wang M, Xu J & Wang S Stretchable transistors and functional circuits for human-integrated electronics. Nat. Electron. 4, 17–29 (2021). [ Google Scholar ] 131. Koo JH et al. A vacuum-deposited polymer dielectric for wafer-scale stretchable electronics. Nat. Electron. 6, 137–145 (2023). [ Google Scholar ] 132. Lee Y et al. Standalone real-time health monitoring patch based on a stretchable organic optoelectronic system. Sci. Adv. 7, eabg9180 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 133. Kang J et al. Ultrathin, solvent-resistant dielectric for monolithic fabrication of low-power, intrinsically stretchable active-matrix electronic skin. Device 2, (2024). [ Google Scholar ] 134. Cui N et al. A photolithographic stretchable transparent electrode for an all-solution-processed fully transparent conformal organic transistor array. J. Mater. Chem. C 7, 5385–5393 (2019). [ Google Scholar ] 135. Yang JC et al. Electronic Skin: Recent Progress and Future Prospects for Skin-Attachable Devices for Health Monitoring, Robotics, and Prosthetics. Adv. Mater. 31, 1904765 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 136. Yang J et al. Masticatory system–inspired microneedle theranostic platform for intelligent and precise diabetic management. Sci. Adv. 8, eabo6900 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 137. Lee W et al. Conformable microneedle pH sensors via the integration of two different siloxane polymers for mapping peripheral artery disease. Sci. Adv. 7, eabi6290 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 138. Zhu Z et al. Blue-ringed octopus-inspired microneedle patch for robust tissue surface adhesion and active injection drug delivery. Sci. Adv. 9, eadh2213 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 139. Sun H, Zheng Y, Shi G, Haick H & Zhang M Wearable Clinic: From Microneedle-Based Sensors to Next-Generation Healthcare Platforms. Small 19, 2207539 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 140. Weibel DB, DiLuzio WR & Whitesides GM Microfabrication meets microbiology. Nat. Rev. Microbiol. 5, 209–218 (2007). [ DOI ] [ PubMed ] [ Google Scholar ] 141. Overholts AC, Granados Razo W & Robb MJ Mechanically gated formation of donor–acceptor Stenhouse adducts enabling mechanochemical multicolour soft lithography. Nat. Chem. 15, 332–338 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 142. Abe Y et al. Porous microneedle-based wearable device for monitoring of transepidermal potential. Biomed. Eng. Adv. 1, 100004 (2021). [ Google Scholar ] 143. Lee H et al. Porous microneedles on a paper for screening test of prediabetes. Med. DEVICES Sens. 3, e10109 (2020). [ Google Scholar ] 144. Wang Y et al. Supramolecular Liquid Crystal Carbon Dots for Solvent-Free Direct Ink Writing. Adv. Mater. 35, 2303680 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 145. Song Y et al. 3D-printed epifluidic electronic skin for machine learning–powered multimodal health surveillance. Sci. Adv. 9, eadi6492 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 146. Lacey SD et al. Extrusion-Based 3D Printing of Hierarchically Porous Advanced Battery Electrodes. Adv. Mater. 30, 1705651 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 147. Zhang L, Wang R, Wang J, Wu L & Zhang X Mechanically robust nanocomposites from screen-printable polymer/graphene nanosheet pastes. Nanoscale 11, 2343–2354 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 148. Li W, Yang S & Shamim A Screen printing of silver nanowires: balancing conductivity with transparency while maintaining flexibility and stretchability. Npj Flex. Electron. 3, 13 (2019). [ Google Scholar ] 149. Quan H et al. Photo-curing 3D printing technique and its challenges. Bioact. Mater. 5, 110–115 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 150. Shin D & Hyun J Silk fibroin microneedles fabricated by digital light processing 3D printing. J. Ind. Eng. Chem. 95, 126–133 (2021). [ Google Scholar ] 151. Melchels FPW, Feijen J & Grijpma DW A review on stereolithography and its applications in biomedical engineering. Biomaterials 31, 6121–6130 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 152. Dong Z et al. 3D printing of inherently nanoporous polymers via polymerization-induced phase separation. Nat. Commun. 12, 247 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 153. Moore DG, Barbera L, Masania K & Studart AR Three-dimensional printing of multicomponent glasses using phase-separating resins. Nat. Mater. 19, 212–217 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 154. Cheng J et al. Centrifugal multimaterial 3D printing of multifunctional heterogeneous objects. Nat. Commun. 13, 7931 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 155. Yin S et al. Wearable and Implantable Soft Robots. Chem. Rev. 124, 11585–11636 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 156. Ge Q et al. 3D printing of highly stretchable hydrogel with diverse UV curable polymers. Sci. Adv. 7, eaba4261 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 157. Dhand AP et al. Additive manufacturing of highly entangled polymer networks. Science 385, 566–572 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 158. Narayan RJ, Doraiswamy A, Chrisey DB & Chichkov BN Medical prototyping using two photon polymerization. Mater. Today 13, 42–48 (2010). [ Google Scholar ] 159. Xing J-F, Zheng M-L & Duan X-M Two-photon polymerization microfabrication of hydrogels: an advanced 3D printing technology for tissue engineering and drug delivery. Chem. Soc. Rev. 44, 5031–5039 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 160. Singh S, Choudhury D, Yu F, Mironov V & Naing MW In situ bioprinting – Bioprinting from benchside to bedside? Acta Biomater. 101, 14–25 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 161. Albanna M et al. In Situ Bioprinting of Autologous Skin Cells Accelerates Wound Healing of Extensive Excisional Full-Thickness Wounds. Sci. Rep. 9, 1856 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 162. Kolan KCR, Semon JA, Bromet B, Day DE & Leu MC Bioprinting with human stem cell-laden alginate-gelatin bioink and bioactive glass for tissue engineering. Int. J. Bioprinting 5, 204 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 163. Zhao W, Hu C & Xu T In vivo bioprinting: Broadening the therapeutic horizon for tissue injuries. Bioact. Mater. 25, 201–222 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 164. Ash C, Dubec M, Donne K & Bashford T Effect of wavelength and beam width on penetration in light-tissue interaction using computational methods. Lasers Med. Sci. 32, 1909–1918 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 165. Chen Y et al. Noninvasive in vivo 3D bioprinting. Sci. Adv. 6, eaba7406 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 166. Urciuolo A et al. Intravital three-dimensional bioprinting. Nat. Biomed. Eng. 4, 901–915 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 167. Kuang X et al. Self-enhancing sono-inks enable deep-penetration acoustic volumetric printing. Science 382, 1148–1155 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 168. Zhang YS et al. 3D extrusion bioprinting. Nat. Rev. Methods Primer 1, 75 (2021). [ Google Scholar ] 169. Gudapati H, Dey M & Ozbolat I A comprehensive review on droplet-based bioprinting: Past, present and future. Biomaterials 102, 20–42 (2016). [ DOI ] [ PubMed ] [ Google Scholar ] 170. Gupta D, Derman ID, Xu C, Huang Y & Ozbolat IT Droplet-based bioprinting. Nat. Rev. Methods Primer 5, 25 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 171. Kim M, Park JJ, Cho C & Ko SH Liquid Metal based Stretchable Room Temperature Soldering Sticker Patch for Stretchable Electronics Integration. Adv. Funct. Mater. 33, 2303286 (2023). [ Google Scholar ] 172. Chung HU et al. Skin-interfaced biosensors for advanced wireless physiological monitoring in neonatal and pediatric intensive-care units. Nat. Med. 26, 418–429 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 173. Zhou S et al. Transcranial volumetric imaging using a conformal ultrasound patch. Nature 629, 810–818 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 174. Li J et al. Vertically stacked skin-like active-matrix display with ultrahigh aperture ratio. Light Sci. Appl. 13, 177 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 175. Lopes PA, Santos BC, de Almeida AT & Tavakoli M Reversible polymer-gel transition for ultra-stretchable chip-integrated circuits through self-soldering and self-coating and self-healing. Nat. Commun. 12, 4666 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 176. Zheng S et al. Pressure-stamped stretchable electronics using a nanofibre membrane containing semi-embedded liquid metal particles. Nat. Electron. 7, 576–585 (2024). [ Google Scholar ] 177. Sunwoo S-H et al. Soft bioelectronics for the management of cardiovascular diseases. Nat. Rev. Bioeng. 2, 8–24 (2023). [ Google Scholar ] 178. Jiang Z et al. A 1.3-micrometre-thick elastic conductor for seamless on-skin and implantable sensors. Nat. Electron. 5, 784–793 (2022). [ Google Scholar ] 179. Zargartalebi H et al. Active-reset protein sensors enable continuous in vivo monitoring of inflammation. Science 386, 1146–1153 (2024). [ DOI ] [ PubMed ] [ Google Scholar ] 180. Ye C et al. A wearable aptamer nanobiosensor for non-invasive female hormone monitoring. Nat. Nanotechnol. 19, 330–337 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 181. Min J et al. Continuous biochemical profiling of the gastrointestinal tract using an integrated smart capsule. Nat. Electron. 8, 844–855 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 182. Tang H et al. Injectable ultrasonic sensor for wireless monitoring of intracranial signals. Nature 630, 84–90 (2024). [ DOI ] [ PubMed ] [ Google Scholar ] 183. Kwak J et al. Monitoring bovine health using wireless mechano-acoustic biophysical sensors. Device 4, 101024 (2026). [ Google Scholar ] 184. Chung HU et al. Skin-interfaced biosensors for advanced wireless physiological monitoring in neonatal and pediatric intensive-care units. Nat. Med. 26, 418–429 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 185. Tehrani F et al. An integrated wearable microneedle array for the continuous monitoring of multiple biomarkers in interstitial fluid. Nat. Biomed. Eng. 6, 1214–1224 (2022). [ DOI ] [ PubMed ] [ Google Scholar ] 186. Liu Y et al. Epidermal mechano-acoustic sensing electronics for cardiovascular diagnostics and human-machine interfaces. Sci. Adv. 2, e1601185 (2016). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 187. Lee K et al. Mechano-acoustic sensing of physiological processes and body motions via a soft wireless device placed at the suprasternal notch. Nat. Biomed. Eng. 4, 148–158 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 188. Cotur Y et al. Stretchable Composite Acoustic Transducer for Wearable Monitoring of Vital Signs. Adv. Funct. Mater. 30, 1910288 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 189. Khodagholy D et al. NeuroGrid: recording action potentials from the surface of the brain. Nat. Neurosci. 18, 310–315 (2015). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 190. Canales A et al. Multifunctional fibers for simultaneous optical, electrical and chemical interrogation of neural circuits in vivo. Nat. Biotechnol. 33, 277–284 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 191. Lu C et al. Flexible and stretchable nanowire-coated fibers for optoelectronic probing of spinal cord circuits. Sci. Adv. 3, e1600955 (2017). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 192. Kim H et al. Bioelectronic Sutures with Electrochemical pH-Sensing for Long-Term Monitoring of the Wound Healing Progress. Adv. Funct. Mater. 34, 2402501 (2024). [ Google Scholar ] 193. Zhao Y et al. Highly Stretchable and Strain-Insensitive Fiber-Based Wearable Electrochemical Biosensor to Monitor Glucose in the Sweat. Anal. Chem. 91, 6569–6576 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 194. Jiang S et al. Spatially expandable fiber-based probes as a multifunctional deep brain interface. Nat. Commun. 11, 6115 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 195. Sahasrabudhe A et al. Multifunctional microelectronic fibers enable wireless modulation of gut and brain neural circuits. Nat. Biotechnol. 42, 892–904 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 196. Wang L et al. Weaving Sensing Fibers into Electrochemical Fabric for Real-Time Health Monitoring. Adv. Funct. Mater. 28, 1804456 (2018). [ Google Scholar ] 197. Wang H et al. Laser Writing of Janus Graphene/Kevlar Textile for Intelligent Protective Clothing. ACS Nano 14, 3219–3226 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 198. Fan W et al. Machine-knitted washable sensor array textile for precise epidermal physiological signal monitoring. Sci. Adv. 6, eaay2840 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 199. Yin D et al. A battery-free nanofluidic intracellular delivery patch for internal organs. Nature 642, 1051–1061 (2025). [ DOI ] [ PubMed ] [ Google Scholar ] 200. Liu Y et al. Soft and elastic hydrogel-based microelectronics for localized low-voltage neuromodulation. Nat. Biomed. Eng. 3, 58–68 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 201. Zhang Y et al. Millimetre-scale bioresorbable optoelectronic systems for electrotherapy. Nature 640, 77–86 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 202. Sahasrabudhe A et al. Multifunctional microelectronic fibers enable wireless modulation of gut and brain neural circuits. Nat. Biotechnol. 42, 892–904 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 203. Lee J et al. Flexible, sticky, and biodegradable wireless device for drug delivery to brain tumors. Nat. Commun. 10, 5205 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 204. Xue H et al. Wearable flexible ultrasound microneedle patch for cancer immunotherapy. Nat. Commun. 16, 2650 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 205. Shirzaei Sani E et al. A stretchable wireless wearable bioelectronic system for multiplexed monitoring and combination treatment of infected chronic wounds. Sci. Adv. 9, eadf7388 (2023). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 206. Jiang Y et al. Wireless, closed-loop, smart bandage with integrated sensors and stimulators for advanced wound care and accelerated healing. Nat. Biotechnol. 41, 652–662 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 207. Li M et al. A wearable and stretchable dual-wavelength LED device for home care of chronic infected wounds. Nat. Commun. 15, 9380 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 208. Minev IR et al. Electronic dura mater for long-term multimodal neural interfaces. Science 347, 159–163 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 209. Oehrn CR et al. Chronic adaptive deep brain stimulation versus conventional stimulation in Parkinson’s disease: a blinded randomized feasibility trial. Nat. Med. 30, 3345–3356 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 210. Wu H et al. Accelerated intestinal wound healing via dual electrostimulation from a soft and biodegradable electronic bandage. Nat. Electron. 7, 299–312 (2024). [ Google Scholar ] 211. Liu Y et al. Soft and elastic hydrogel-based microelectronics for localized low-voltage neuromodulation. Nat. Biomed. Eng. 3, 58–68 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 212. Zheng Z, Zhu R, Peng I, Xu Z & Jiang Y Wearable and implantable biosensors: mechanisms and applications in closed-loop therapeutic systems. J. Mater. Chem. B 12, 8577–8604 (2024). [ DOI ] [ PubMed ] [ Google Scholar ] 213. Kilgard MP et al. Closed-loop vagus nerve stimulation aids recovery from spinal cord injury. Nature 643, 1030–1036 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 214. Ferrero JJ et al. Closed-loop electrical stimulation prevents focal epilepsy progression and long-term memory impairment. Nat. Neurosci. 28, 1753–1762 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 215. Ouyang W et al. A wireless and battery-less implant for multimodal closed-loop neuromodulation in small animals. Nat. Biomed. Eng. 7, 1252–1269 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 216. Wang W et al. Neuromorphic sensorimotor loop embodied by monolithically integrated, low-voltage, soft e-skin. Science 380, 735–742 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 217. Mickle AD et al. A wireless closed-loop system for optogenetic peripheral neuromodulation. Nature 565, 361–365 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 218. Xu C, Mao G, Williamson R & Whiting D Delayed intracerebral hemorrhage: A rare complication of deep brain stimulation surgery. Interdiscip. Neurosurg. 14, 135–138 (2018). [ Google Scholar ] 219. Feiner R & Dvir T Tissue–electronics interfaces: from implantable devices to engineered tissues. Nat. Rev. Mater. 3, 17076 (2017). [ Google Scholar ] 220. Khatsenko K, Khin Y & Maibach H Allergic Contact Dermatitis to Components of Wearable Adhesive Health Devices. Dermatitis 31, 283–286 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 221. Baek J-Y, An J-H, Choi J-M, Park K-S & Lee S-H Flexible polymeric dry electrodes for the long-term monitoring of ECG. Sens. Actuators Phys. 143, 423–429 (2008). [ Google Scholar ] 222. Li Y, Li N, De Oliveira N & Wang S Implantable bioelectronics toward long-term stability and sustainability. Matter 4, 1125–1141 (2021). [ Google Scholar ] 223. Liu J et al. Syringe-injectable electronics. Nat. Nanotechnol. 10, 629–636 (2015). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 224. Seymour JP & Kipke DR Neural probe design for reduced tissue encapsulation in CNS. Biomaterials 28, 3594–3607 (2007). [ DOI ] [ PubMed ] [ Google Scholar ] 225. Ho T-C et al. Hydrogels: Properties and Applications in Biomedicine. Molecules 27, 2902 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 226. Hydrogels. in Biomaterials Science 153–166 (Academic Press, 2020). doi: 10.1016/B978-0-12-816137-1.00014-3. [ DOI ] [ Google Scholar ] 227. Yuk H, Lu B & Zhao X Hydrogel bioelectronics. Chem. Soc. Rev. 48, 1642–1667 (2019). [ DOI ] [ PubMed ] [ Google Scholar ] 228. Hui Y et al. Three-dimensional printing of soft hydrogel electronics. Nat. Electron. 5, 893–903 (2022). [ Google Scholar ] 229. Ohm Y et al. An electrically conductive silver–polyacrylamide–alginate hydrogel composite for soft electronics. Nat. Electron. 4, 185–192 (2021). [ Google Scholar ] 230. Mariani E, Lisignoli G, Borzì RM & Pulsatelli L Biomaterials: Foreign Bodies or Tuners for the Immune Response? Int. J. Mol. Sci. 20, 636 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 231. Veiseh O & Vegas AJ Domesticating the foreign body response: Recent advances and applications. Adv. Drug Deliv. Rev. 144, 148–161 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 232. Wang MO et al. Evaluation of the In Vitro Cytotoxicity of Cross-Linked Biomaterials. Biomacromolecules 14, 1321–1329 (2013). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 233. Imlimthan S et al. Systematic in vitro biocompatibility studies of multimodal cellulose nanocrystal and lignin nanoparticles. J. Biomed. Mater. Res. A 108, 770–783 (2020). [ DOI ] [ PubMed ] [ Google Scholar ] 234. Witherel CE, Abebayehu D, Barker TH & Spiller KL Macrophage and Fibroblast Interactions in Biomaterial-Mediated Fibrosis. Adv. Healthc. Mater. 8, 1801451 (2019). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 235. Kutner N, Kunduru KR, Rizik L & Farah S Recent Advances for Improving Functionality, Biocompatibility, and Longevity of Implantable Medical Devices and Deliverable Drug Delivery Systems. Adv. Funct. Mater. 31, 2010929 (2021). [ Google Scholar ] 236. Jiang S & Cao Z Ultralow-Fouling, Functionalizable, and Hydrolyzable Zwitterionic Materials and Their Derivatives for Biological Applications. Adv. Mater. 22, 920–932 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 237. Knop K, Hoogenboom R, Fischer D & Schubert US Poly(ethylene glycol) in Drug Delivery: Pros and Cons as Well as Potential Alternatives. Angew. Chem. Int. Ed. 49, 6288–6308 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 238. Xie X et al. Reduction of measurement noise in a continuous glucose monitor by coating the sensor with a zwitterionic polymer. Nat. Biomed. Eng. 2, 894–906 (2018). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 239. He W, McConnell GC & Bellamkonda RV Nanoscale laminin coating modulates cortical scarring response around implanted silicon microelectrode arrays. J. Neural Eng. 3, 316 (2006). [ DOI ] [ PubMed ] [ Google Scholar ] 240. Zaveri TD et al. Contributions of surface topography and cytotoxicity to the macrophage response to zinc oxide nanorods. Biomaterials 31, 2999–3007 (2010). [ DOI ] [ PubMed ] [ Google Scholar ] 241. Capuani S, Malgir G, Chua CYX & Grattoni A Advanced strategies to thwart foreign body response to implantable devices. Bioeng. Transl. Med. 7, e10300 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 242. Robotti F et al. A micron-scale surface topography design reducing cell adhesion to implanted materials. Sci. Rep. 8, 10887 (2018). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 243. Booth MA et al. Polycrystalline diamond coating on 3D printed titanium scaffolds: Surface characterisation and foreign body response. Mater. Sci. Eng. C 130, 112467 (2021). [ DOI ] [ PubMed ] [ Google Scholar ] 244. Shelton RM, Rasmussen AC & Davies JE Protein adsorption at the interface between charged polymer substrata and migrating osteoblasts. Biomaterials 9, 24–29 (1988). [ DOI ] [ PubMed ] [ Google Scholar ] 245. Proteins at Interfaces II, Copyright, 1995 Advisory Board, Foreword. in Proteins at Interfaces II vol. 602 i–iv (American Chemical Society, 1995). [ Google Scholar ] 246. Haynes CA & Norde W Globular proteins at solid/liquid interfaces. Colloids Surf. B Biointerfaces 2, 517–566 (1994). [ Google Scholar ] 247. Wilson CJ, Clegg RE, Leavesley DI & Pearcy MJ Mediation of Biomaterial–Cell Interactions by Adsorbed Proteins: A Review. Tissue Eng. 11, 1–18 (2005). [ DOI ] [ PubMed ] [ Google Scholar ] 248. Hunt JA, Flanagan BF, McLaughlin PJ, Strickland I & Williams DF Effect of biomaterial surface charge on the inflammatory reponse: Evaluation of cellular infiltration and TNFα production. J. Biomed. Mater. Res. 31, 139–144 (1996). [ DOI ] [ PubMed ] [ Google Scholar ] 249. Lopez-Silva TL et al. Chemical functionality of multidomain peptide hydrogels governs early host immune response. Biomaterials 231, 119667 (2020). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 250. Zhang J et al. Novel Balanced Charged Alginate/PEI Polyelectrolyte Hydrogel that Resists Foreign-Body Reaction. ACS Appl. Mater. Interfaces 10, 6879–6886 (2018). [ DOI ] [ PubMed ] [ Google Scholar ] 251. Hernandez JL et al. Effect of tissue microenvironment on fibrous capsule formation to biomaterial-coated implants. Biomaterials 273, 120806 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 252. Li Y, Li N, Oliveira ND & Wang S Implantable bioelectronics toward long-term stability and sustainability. Matter 4, 1125–1141 (2021). [ Google Scholar ] 253. Shi D, Dhawan V & Cui XT Bio-integrative design of the neural tissue-device interface. Curr. Opin. Biotechnol. 72, 54–61 (2021). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 254. Hu M, Liang C & Wang D Implantable bioelectrodes: challenges, strategies, and future directions. Biomater. Sci. 12, 270–287 (2024). [ DOI ] [ PubMed ] [ Google Scholar ] 255. Villa J, Cury J, Kessler L, Tan X & Richter C-P Enhancing biocompatibility of the brain-machine interface: A review. Bioact. Mater. 42, 531–549 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 256. Kozai TDY, Jaquins-Gerstl AS, Vazquez AL, Michael AC & Cui XT Brain Tissue Responses to Neural Implants Impact Signal Sensitivity and Intervention Strategies. ACS Chem. Neurosci. 6, 48–67 (2015). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 257. Mariello M Reliability and stability of Bioelectronic Medicine: a critical and pedagogical perspective. Bioelectron. Med. 11, 16 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 258. Zhang Y et al. Advances in Bioresorbable Materials and Electronics. Chem. Rev. 123, 11722–11773 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 259. Hwang S-W et al. A Physically Transient Form of Silicon Electronics. Science 337, 1640–1644 (2012). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 260. Park JY et al. Lessons Learned and Challenges Ahead in the Translation of Implantable Microscale Sensors and Actuators. Annu. Rev. Biomed. Eng. 27, 211–233 (2025). [ DOI ] [ PubMed ] [ Google Scholar ] 261. Submission and Review of Sterility Information in Premarket Notification (510(k)) Submissions for Devices Labeled as Sterile. US Food Drug Adm. (2024). [ Google Scholar ] 262. Shelf Life of Medical Devices, Guidance Document. US Food Drug Adm. (1991). [ Google Scholar ] 263. Biological evaluation of medical devices—Part 1: Evaluation and testing within a risk management process. US Food Drug Adm. http://array.aami.org/doi/10.2345/9781570207556.ch1 (2023). [ Google Scholar ] 264. Chemical Analysis for Biocompatibility Assessment of Medical Devices. US Food Drug Adm. (2024). [ Google Scholar ] 265. Jiang Z et al. A 1.3-micrometre-thick elastic conductor for seamless on-skin and implantable sensors. Nat. Electron. 5, 784–793 (2022). [ Google Scholar ] 266. Khodagholy D et al. NeuroGrid: recording action potentials from the surface of the brain. Nat. Neurosci. 18, 310–315 (2015). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 267. Min J et al. Continuous biochemical profiling of the gastrointestinal tract using an integrated smart capsule. Nat. Electron. 8, 844–855 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 268. Minev IR et al. Electronic dura mater for long-term multimodal neural interfaces. Science 347, 159–163 (2015). [ DOI ] [ PubMed ] [ Google Scholar ] 269. Huang H et al. A gastric retentive robotic capsule enables emergency-prepared and responsive oral drug delivery in canine models. Sci. Adv. 11, eady6745 (2025). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 270. Sun Z et al. A bioabsorbable mechanoelectric fiber as electrical stimulation suture. Nat. Commun. 15, 8462 (2024). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] 271. Song M et al. Full textile-based body-coupled electrical stimulation for wireless, battery-free, and wearable bioelectronics. Npj Flex. Electron. 8, 76 (2024). [ Google Scholar ] 272. Wang W et al. Neuromorphic sensorimotor loop embodied by monolithically integrated, low-voltage, soft e-skin. Science 380, 735–742 (2023). [ DOI ] [ PubMed ] [ Google Scholar ] 273. Lin R et al. Digitally-embroidered liquid metal electronic textiles for wearable wireless systems. Nat. Commun. 13, 2190 (2022). [ DOI ] [ PMC free article ] [ PubMed ] [ Google Scholar ] Associated Data This section collects any data citations, data availability statements, or supplementary materials included in this article. Data Availability Statement No primary research data were generated or analyzed as part of this Review. 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