Advanced Hub Main Navigation Menu
Intrinsically Soft Implantable Electronics for Long-term Biosensing Applications
[Correction added on 26 May 2025, after first online publication: the third author's name was corrected in this version.]
Abstract
Implantable biosensors play a critical role in healthcare and medical research by enabling real-time monitoring of physiological signals with high precion. Compared to non-invasive biosensors, implantable biosensors offer superior fidelity by minimizing external noise and ensuring direct contact with target tissues. However, conventional implantable biosensors, often composed of intrinsically rigid materials such as silicon and metals, suffer from mechanical mismatches with soft biological tissues, leading to inflammatory responses, fibrotic encapsulation, and long-term instability. To address these challenges, recent advances have focused on the development of intrinsically soft materials, which leverage soft and stretchable materials to achieve long-term biocompatibility and seamless tissue integreation. These materials have shown significant promise in neural interfaces, cardiac monitoring, and soft bioelectrodes for cronic sensing and stimulation. This review provides a comprehensive overview of these emerging biosensors, starting with a discussion of the limitations of conventional implantable biosensors. It then examines key intrinsically soft materials, including encapsulation matrices and stretchable conductors, and explores strategies for minimally invasive implantation, chronic fixation, and biocompatibility enhancement. Additionally, specific application cases are highlighted to demonstrate their practical utility. Finally, remaining challenges and future research opportunities are discussed to guide the next generation of intrinsically soft implantable biosensors toward clinical translation.
1 Introduction
Monitoring biological signals plays a pivotal role in understanding the mechanisms of living organisms and applying this knowledge in medical contexts.[1] Abnormalities in these signals often serve as indicators of pathological conditions, necessitating precise detection tools. Biosensors have been developed to collect a wide range of biological signals, including electrophysiological, chemical, physical, and mechanical data, for both scientific research and clinical diagnostics. Non-invasive biosensors, such as wearable devices and lab-on-a-chip platforms, have revolutionized real-time monitoring and personalized medicine through their wearability, portability, and ease of use.[2, 3] However, these devices frequently encounter challenges in signal accuracy due to artifact and noise during data collection, which compromise their signal-to-noise ratio (SNR).
The primary limitation of non-invasive biosensors lies in their inability to directly interface with the target organ. Biological signals must traverse multiple tissue layers, such as skin, muscle, and bone, leading to significant signal attenuation and interference. Consequently, these sensors often lack spatiotemporal specificity, rendering them inadequate for applications requiring precise, localized data collection.
Implantable biosensors overcome these limitations by directly interfacing with target tissues, enabling high-fidelity signal acquisition. However, traditional implantable devices, composed of rigid materials such as silicon and metals, face significant hurdles due to the mismatch in their mechanical properties from those of soft biological tissues (hereinafter referred to as “mechanical mismatch”).[4, 5] The mechanical mismatch provokes inflammatory responses, leading to the formation of fibrous encapsulation that degrades device performance over time.[6] Recent innovations have focused on ultrathin flexible sensors, incorporating stress-relieving designs such as serpentine and buckling structures to mitigate these mismatches. While effective to some extent, these approaches remain limited by the intrinsic rigidity of the materials, resulting in mechanical fatigue and fabrication complexity.
To address these issues, recent research has focused on intrinsically soft materials that can match the mechanical properties of the biological tissues, reducing foreign body responses and enhancing long-term stability.[7, 8] Unlike conventional flexible bioelectronics that rely on ultrathin rigid materials or stress-absorbing designs, intrinsically soft biosensors offer an alternative approach that fundamentally mitigates mechanical mismatch issues. Various classes of intrinsically soft materials, including elastomers, hydrogels, and liquid metals, have been explored for their potential to enhance biocompatibility and durability in long-term implantation. This review aims to provide a comprehensive discussion on the advances in intrinsically soft implantable biosensors, addressing the following key questions: 1) What are the primary complications associated with conventional implantable biosensors, and how do intrinsically soft materials overcome these limitations? 2) What are the most promising material candidates for encapsulation and conductive components in soft bioelectronics? 3) What strategies have been developed for minimally invasive implantation, long-term fixation, and chronic biocompatibility? 4) How have these materials been successfully applied in neural, cardiac, and biochemical sensing applications? 5) What are the remaining technical challenges and future research directions for clinical translation?
By systematically addressing these questions, this review aims to offer valuable insights into the design, fabrication, and application of intrinsically soft biosensors, providing guidance for researchers in both academia and industry toward the next generation of bioelectronic systems (Figure 1).[9, 10]
2 Acute and Chronic Inflammatory Responses Upon Implantation
Implantable biosensors hold significant promise for precise, spatiotemporal collection of high-quality biological signals. However, the introduction of foreign materials into biological tissues inevitably triggers complex inflammatory responses. These responses, while part of the body's defense mechanisms, can negatively impact the performance and longevity of implantable devices. The inflammatory responses can be broadly classified into acute and chronic phases, each characterized by distinct cellular and molecular processes.
2.1 Acute Inflammation
Acute inflammation occurs immediately following implantation and represents the body's rapid response to injury or foreign materials. This phase is characterized by vascular changes, immune cell recruitment, and the activation of inflammatory mediators. Vasodilation and increased capillary permeability facilitate the migration of immune cells and proteins to the site of injury. Neutrophils are the first responders, releasing reactive oxygen species and proteolytic enzymes to clear debris and pathogens. However, these actions may also damage surrounding tissues, exacerbating inflammation. Within 24–48 h, monocytes infiltrate the site and differentiate into macrophages, which play dual roles: clearing debris and orchestrating tissue repair by secreting cytokines and growth factors.[11, 12] The resolution of acute inflammation is critical for initiating tissue healing and avoiding progression to chronic inflammation.
2.2 Chronic Inflammation
Chronic inflammation arises when the acute phase fails to resolve, often due to the persistent presence of foreign materials or prolonged tissue damage. This phase involves sustained immune activation, characterized by macrophage accumulation, fibroblast proliferation, and extracellular matrix deposition. Over time, chronic inflammation can lead to fibrosis and the formation of a dense, collagen-rich capsule around the implant. This encapsulation isolates the device from surrounding tissues, impairing its functionality by reducing sensitivity, increasing signal interference, and compromising biocompatibility (Figure 2a).[13] Furthermore, the chronic inflammatory environment can accelerate material degradation, shortening the device's operational lifespan. Effective strategies to mitigate chronic inflammation include using biocompatible materials, minimizing mechanical mismatch, and functionalizing device surfaces to reduce immune activation. In case of rigid materials, mechanical mismatch of the material and implant severely exacerbate the acute and chronic inflammation (Figure 2b). Understanding the mechanisms underlying acute and chronic inflammation is vital for designing implantable biosensors that minimize adverse biological responses while ensuring reliable and durable performance. The following sections will explore material innovations and strategies to match the mechanical modulus of biosensors within physiological environments (Figure 2c).[14]
3 Flexible Bioelectrodes Comprising Intrinsically Rigid Materials
Flexible bioelectrodes are critical for biomedical applications such as neural interfaces, cardiac monitoring, and wearable health devices. Conventional bioelectrodes, often made from rigid materials like silicon, metals, and plastics, provide excellent electrical properties but lack the flexibility needed for effective integration with soft, dynamic tissues. To overcome these limitations, researchers have developed methods to enhance the flexibility and stretchability of these materials, enabling better conformability to biological surfaces.[15] This section examines strategies for improving bioelectrode flexibility, focusing on ultrathin structures and advanced design techniques for stretchability. It also addresses the challenges associated with using intrinsically rigid materials, including mechanical mismatch with soft tissues and potential long-term degradation in performance, emphasizing the need for continued innovation in this field.
3.1 Ultrathin Structure for Flexibility
A fundamental strategy for enhancing the flexibility of intrinsically rigid materials, such as silicon and metals, is to reduce their thickness to the nanoscale. Ultrathin structures allow these materials to bend and flex without breaking, making them ideal for integration into flexible bioelectronic devices.[16, 17] The reduction in thickness minimizes the deformation disparity between the inner and outer surfaces during bending, significantly reducing mechanical stress and enhancing durability. Bulk forms of silicon and metals are traditionally rigid and brittle, but when reduced to thicknesses below 100 nanometers, their bending stiffness decreases dramatically.[18] This property allows these materials to conform to complex and dynamic biological surfaces without generating significant stress. For example, ultrathin silicon or metallic films can bend to tight radii of curvature, maintaining their structural integrity and functionality under repeated deformation.
3.2 Structural Design for Stretchability
Enhancing stretchability in intrinsically rigid materials is critical for applications requiring adaptation to dynamic environments, such as implantable devices interfacing with soft biological tissues. While reducing material thickness to nanoscale levels improves flexibility, additional structural innovations are required to achieve stretchability.[19] Advanced design techniques enable these materials to absorb and redistribute mechanical stress effectively, maintaining their integrity under repeated deformation.
3.2.1 In-Plane Serpentine Structure
In-plane serpentine structures involve arranging conductive or structural materials into wavy, snake-like patterns (Figures 3a,b).[20] This geometry allows materials to elongate and compress without fracturing, distributing stress uniformly along the entire trace rather than concentrating it at specific points. Strain-to-failure rates in serpentine designs range widely, depending on parameters like ribbon width and arc radius. For example, Ji et al. developed serpentine-patterned parylene C electrodes for electrocorticography (ECoG) applications, achieving stretchability up to 50% and improving long-term tissue compatibility (Figure 3c).[21] This approach underscores the potential of serpentine designs for bioelectronics requiring high mechanical compliance.
3.2.2 Out-of-Plane Buckling Structure
Out-of-plane buckling structures exploit the tendency of thin films to form periodic undulations under compressive forces (Figure 3d).[18] These buckled patterns can accommodate significant strain perpendicular to the buckling direction (Figures 3e and f).[22] Siddiqui et al. demonstrated piezoelectric nanogenerators with omnidirectional stretchability using buckled graphite electrodes.[23] These devices maintained electrical stability over 9000 strain cycles, highlighting their durability and suitability for wearable and implantable devices.
3.2.3 Helical Structure
Helical configurations utilize coiled designs, allowing rigid materials to stretch and twist while maintaining structural integrity (Figures 3g and h).[24] Stress is uniformly distributed along the coil, enabling high levels of deformation without failure. Helical designs are particularly advantageous for neural and cardiac interfaces, where devices must conform to soft, curved biological tissues (Figure 3i).[25]
3.2.4 Kirigami and Origami
Kirigami and origami techniques utilize cutting and folding to create stretchable, adaptable structures. Kirigami designs excel in forming intricate, expandable patterns, while origami-based folding produces three-dimensional, compact structures capable of significant deformation (Figures 3j–l).[26] For instance, Li et al. integrated kirigami structures into triboelectric nanogenerators, enhancing mechanical adaptability for joint therapy. Despite their potential, kirigami designs are limited in compressibility, requiring additional modifications to expand their applicability. By combining these advanced structural techniques with ultrathin materials, researchers are advancing the functionality and resilience of stretchable bioelectronics. These innovations provide the foundation for seamless integration with soft tissues while addressing the challenges posed by the inherent rigidity of conventional materials.
3.3 Complications of the Deformable Electronics Using Intrinsically Rigid Materials
Despite advancements in designing stretchable electronics from intrinsically rigid materials, several inherent limitations hinder their applicability, particularly in soft implantable biosensors. These challenges stem from the complexity of fabrication processes, spatial constraints, mechanical durability, and encapsulation compatibility.
Fabrication of intricate designs, such as serpentine, buckling, or origami structures, requires sophisticated and costly techniques, including photolithography and etching. These processes increase production time and expenses, making scalability for widespread biomedical applications challenging (Figure 3m).[20, 27] Additionally, the precision needed to achieve the required stretchability adds another layer of complexity.
Stretchable designs demand extra space for deformation, such as the reserved lengths in serpentine or buckling patterns, which reduces the functional density of devices (Figure 3n).[28] This limitation makes it difficult to develop compact, high-performance systems essential for implantation, particularly in confined biological environments.
The intrinsic rigidity of materials also leads to fatigue and potential failure under repeated mechanical stress.[29] While stretchable designs help distribute strain, they cannot entirely prevent cyclic loading-induced degradation (Figure 3o). Such wear compromises the long-term durability of devices, an issue critical for applications requiring continuous operation in dynamic conditions.[30, 31]
Moreover, rigid materials are often incompatible with soft encapsulation materials, resulting in unreliable sealing (Figure 3p).[32] Mismatched mechanical properties between rigid components and flexible encapsulation layers can cause delamination, cracking, or leakage. This undermines device integrity, impairs biocompatibility, and increases the risk of device failure in physiological environments.
Addressing these challenges necessitates innovative approaches in material selection and fabrication techniques. The following sections will delve into the development of intrinsically soft materials and fabrication strategies that aim to overcome these limitations while maintaining the required functionality and durability for implantable biosensors.
4 Intrinsically Soft Matrices
In the realm of soft and implantable biosensors, the selection of substrate and encapsulating materials is paramount to optimizing device performance, biocompatibility, and long-term stability. These materials must adapt to dynamic and irregular biological tissue surfaces while ensuring mechanical robustness and protecting embedded sensor components. Additionally, they should exhibit biocompatibility to reduce immune responses and sufficient stability to function reliably in physiological environments. This chapter delves into various types of soft matrices, including physically crosslinked elastomers (Figure 4a), chemically crosslinked elastomers (Figure 4b), and hydrogels, emphasizing their role in soft biosensor applications.
4.1 Physically Cross-Linked Elastomers
Physically crosslinked elastomers derive their elasticity from non-covalent interactions, such as phase separation or polymer chain entanglements (Figure 4a).[33, 34] These materials are widely utilized for their excellent mechanical flexibility, ease of processing, and adaptability to biomedical applications. By tuning the ratio of hard and soft segments in the polymer structure, their mechanical and functional properties can be tailored to specific needs.
4.1.1 Thermoplastic Elastomers
Thermoplastic polyurethanes (TPUs) are renowned for their high elasticity, toughness, and biocompatibility (Figure 4c).[35, 36] Their segmented block copolymer structure features soft segments providing flexibility and hard segments ensuring mechanical strength. TPUs find applications in cardiac patches, vascular grafts, and wearable sensors, where materials must endure repetitive deformation.[37] However, TPU's susceptibility to hydrolysis in humid conditions necessitates chemical modifications with hydrophobic additives to enhance durability.
4.1.2 Block Copolymers
Block copolymers, such as styrene-butadiene-styrene (SBS) and styrene-ethylene-butylene-styrene (SEBS),[38] combine alternating rigid and flexible polymer blocks, yielding elastomers with superior mechanical stability and elasticity (Figure 4d).[39, 40] SEBS, known for its thermal and oxidative resistance, is extensively used in stretchable electronics and biosensor encapsulation.[39] SBS's rubber-like elasticity at room temperature and thermoplastic processability make it ideal for flexible device fabrication. Both materials deliver high tensile strength and elongation, ensuring durable and conformable sensor designs.
4.2 Chemically Cross-Linked Elastomers
Chemically crosslinked elastomers form covalent bonds between polymer chains, creating robust three-dimensional networks (Figure 4b).[33, 41] These materials offer superior mechanical strength, thermal stability, and solvent resistance, making them suitable for demanding biomedical applications.
4.2.1 Silicone
Silicone elastomers, particularly polydimethylsiloxane (PDMS), are among the most extensively used materials in biomedical engineering due to their flexibility, biocompatibility, and chemical inertness (Figure 4e).[42, 43] PDMS finds applications in microfluidic devices, neural probes, and epidermal electronics.[42] By adjusting the curing agent ratio, the mechanical properties of PDMS can be finely tuned. However, excessive crosslinking may introduce residual agents that negatively affect material stability. Ecoflex, a softer silicone variant, is preferred in applications requiring extreme elasticity, such as soft robotics and dynamic implants.
4.2.2 Photocrosslinking Polymers (PUA)
Polyurethane acrylates (PUAs) combine the mechanical resilience of polyurethane with the precise patterning capabilities of acrylates (Figure 4f).[44-46] Cured using UV light, PUAs enable the fabrication of microscale features essential for biosensor applications. While offering excellent mechanical and chemical properties, balancing crosslinking density is crucial to maintain flexibility and biocompatibility for implantable uses.
4.3 Hydrogels
Hydrogels, characterized by their high water content (70–90%) and softness, mimic the mechanical properties of biological tissues.[14] Their biocompatibility and tunable characteristics make them indispensable in applications such as glucose monitoring, wound healing, and drug delivery.[47, 48] Hydrogels facilitate molecular diffusion, enabling biosensors to interact seamlessly with their environment.[6, 49] Chemically crosslinked hydrogels provide structural stability under physiological conditions, while physically crosslinked hydrogels exhibit dynamic sol-gel transitions for responsive capabilities.
Hydrogels’ mechanical and electrical weaknesses can be mitigated by reinforcement with nanoparticles, nanofibers, or additional crosslinking agents.[50, 51] This reinforcement enhances their structural integrity, ensuring consistent functionality. However, challenges such as dehydration—which can alter their properties—necessitate precise design strategies to maintain performance over extended periods.[52]
A comprehensive explanation about the pros and cons of the elastomeric materials introduced above are demonstrated in Table 1.
| Materials | Advantages | Limitations | |
|---|---|---|---|
| Physically crosslinked elastomers | Thermoplastic polyurethane (TPU) |
|
|
| Block-copolymers |
|
|
|
| Chemically crosslinked elastomers | Polydimethylsiloxane (PDMS) |
|
|
| Ecoflex |
|
|
|
| Polyurethane Acrylates (PUA) |
|
|
|
| Hydrogels | Physically Crosslinked Hydrogels |
|
|
| Chemically Crosslinked Hydrogels |
|
|
|
4.4 Conductive Polymers and Hydrogels
Conductive polymers and hydrogels integrate electrical conductivity with mechanical softness, broadening their utility in flexible and stretchable electronics. Their unique properties make them particularly valuable for biosensors interfacing with biological systems.
4.4.1 Conductive Polymers
Poly(3,4-Ethylenedioxythiophene) (PEDOT)
PEDOT is a highly conductive polymer known for its excellent stability, transparency, and flexibility. In its doped form, PEDOT:PSS (poly(styrene sulfonate)) is water-soluble and easier to process (Figure 5a).[53] The conductivity of PEDOT:PSS arises from the conjugated structure of PEDOT, which allows for the delocalization of electrons along its polymer backbone. However, PSS, which acts as the counterion for PEDOT, is an insulating polymer that facilitates water solubility but can limit the overall conductivity of the material. In PEDOT:PSS, PEDOT exists as positively charged chains, and PSS as the negatively charged sulfonate groups that stabilize the PEDOT chains in the solution. While this structure is beneficial for processability, the insulating nature of PSS reduces the efficiency of electron transport within the material.[54] PEDOT:PSS films are transparent, flexible, and exhibit high electrical conductivity, making them suitable for various electronic and optoelectronic applications. In the context of biosensors, PEDOT:PSS serves as a conductive coating for electrodes, enhancing their electrical properties while maintaining flexibility. It is also used in stretchable sensors and wearable devices, where its combination of conductivity and flexibility is highly desirable.
For example, Tan et al. developed a self-adhesive conductive polymer composite comprising PEDOT:PSS dispersed in poly(vinyl alcohol) crosslinked by glutaraldehyde.[55] The composite possess high mechanical flexibility (56.1 kPa), stretchability (≈700%), and electrical conductivity (37 S cm−1) simultaneously. Moreover, the PEDOT:PSS composite demonstrated exceptional transparency (95%) and strong adhesion on various surfaces, including polyimide (PI) film, skin, and internal organs (Figure 5b). To overcome the limitation on low conductivity and poor mechanical modulus, various post-treatment methods are applied. Acid doping, for instance, involves the use of strong acids (e.g., sulfuric acid) to remove excess PSS and improve the conductive pathways by reorganizing the PEDOT chains.[56] Solvent annealing, using solvents such as ethylene glycol or dimethyl sulfoxide (DMSO), helps to phase-separate PEDOT from PSS, further reducing the insulating effect of PSS and increasing the overall conductivity of the film.[57]
In another example, Wang et al. presented a highly stretchable, transparent, and conductive polymer based on PEDOT:PSS enhanced with ionic additives.[58] The resulting PEDOT:PSS demonstrated that such additives significantly improve both conductivity and mechanical properties, enabling the polymer to maintain high electrical performance even under large strains (4100 S cm−1 under 100%) (Figure 3c). The mechanism involves creating a conductive nanofibrillar network within a soft matrix by weakening the electrostatic interaction between PEDOT and PSS, thus allowing better crystallinity and stretchability. Recently, Oh et al. added ionic liquid to PEDOT:PSS to enhance electrical conductivity (286 S cm−1) and mechanical stretchability (90% strain) simultaneously.[59] The ionic liquids helped maintain the mobility of charge carriers and improved the mechanical properties due to their homogeneous distribution, forming hydrogen-bonded networks of densely packed PEDOT colloids, which could also be used as a printable ink (Figure 5d).
Polyaniline (PANI)
Polyaniline (PANI) is another conductive polymer known for its tunable conductivity, which can be adjusted by varying its oxidation and protonation state (Figure 5e). PANI exists in several oxidation states, including leucoemeraldine (fully reduced, insulating), emeraldine (half-oxidized, conductive), and pernigraniline (fully oxidized, insulating).[60] Among these, the emeraldine salt form is the most conductive and is widely used in electronic applications. The conductivity of PANI arises primarily from its ability to exist in multiple oxidation states, combined with protonation by acids.[61] In its emeraldine salt form, PANI's polymer backbone contains alternating amine (─NH─) and imine (─N═) groups. When protonated by acids (e.g., hydrochloric acid), the imine groups become positively charged, allowing the formation of delocalized polarons (charged carriers) along the polymer chain. This polaron formation is what enables electrical conductivity in PANI. The level of protonation can be tuned, allowing precise control over the material's conductivity, which makes PANI particularly versatile for different applications. In contrast, the fully reduced leucoemeraldine state and the fully oxidized pernigraniline state are non-conductive because these forms do not support delocalized charge carriers.[62] Thus, PANI's conductivity is highest in its half-oxidized, protonated emeraldine salt form, where it acts as a conductive polymer suitable for various electronic and biosensing applications (Figure 5f).[63, 64]
Polypyrrole (PPy)
Polypyrrole (PPy) is a conductive polymer widely studied for its electrical conductivity, biocompatibility, and ease of synthesis (Figure 5g). It is typically synthesized through oxidative polymerization, resulting in a black, electrically conductive material. The conductivity of PPy arises from its conjugated polymer backbone, which allows for the delocalization of π-electrons along the polymer chains.[65] This delocalization enables the movement of charge carriers, such as polarons and bipolarons, which are generated through doping.[66]
In its neutral state, PPy is insulating, but upon oxidation (doping), it becomes conductive as charge carriers are generated along the polymer chain.[67] The doping process typically involves adding counterions—often small anions, such as chloride or perchlorate—which balance the positive charges generated during polymerization, thus making PPy conductive. These counterions stabilize the charged sites on the polymer chain and enable efficient charge transport.[68]
PPy is stable in both air and water, making it suitable for various biomedical applications. It is commonly used in biosensors, neural interfaces, and actuators. In biosensors, PPy can act as a transducer material, converting biological signals into electrical signals. For instance, PPy-based coatings have been used to enhance the sensitivity and selectivity of glucose sensors by providing a conductive interface for enzyme immobilization and signal transduction. In neural interfaces, PPy is employed as a conductive coating on electrodes to improve the interface between the electrode and neural tissue, enhancing signal transduction and minimizing tissue reaction. He et al. demonstrated a conductive nanofiber hydrogel with a high mechanical strength (1.6 MPa and 55% elongation) and electronic conductivity (Figure 5h).[69] The PPy infiltrated in the aramid nanofiber structure showed an ultralow percolation threshold (≈1 wt.%). At higher PPy concentrations, the electrical conductivity reached 80 S cm−1.
4.4.2 Conductive Hydrogels
Conductive hydrogels have garnered significant interest as materials for implantable biosensors due to their unique combination of electrical conductivity and soft, tissue-like mechanical properties. Hydrogels are water-rich, three-dimensional polymer networks that can mimic the softness and flexibility of biological tissues. When these hydrogels are made conductive—either through ionic conduction or by integrating conductive polymers—they become highly suitable for bioelectronic applications.
Ionically Conductive Hydrogels
Ionically conductive hydrogels achieve conductivity through the movement of free ions within their water-rich network (Figure 5i). Unlike traditional conductive materials that rely on electron transport, these hydrogels facilitate ionic conduction, making them ideal for biosensors that interface with biological systems, where ionic signaling is crucial. The electrical conductivity of ionically conductive hydrogels stems from the dissociation of ionic species, such as salts, acids, or polyelectrolytes, within the hydrogel matrix. When an electric field is applied, these ions move, creating a pathway for ionic current. The high-water content of the hydrogel enhances ion mobility, mimicking the way biological tissues transmit signals through ion flow.[70, 71]
Ionically conductive hydrogels are highly biocompatible due to their high-water content and similarity to natural tissues, minimizing immune response and ensuring long-term stability.[72] Their mechanical properties also allow them to conform to soft tissues without causing damage, making them suitable for long-term implants in dynamic biological environments. However, their ionic conduction mechanism often results in lower conductivity compared to electron-conducting materials. While sufficient for some biosensing applications, this limitation may restrict their use in applications that require higher electrical performance. Over time, ionically conductive hydrogels may degrade or swell, leading to changes in their conductivity and mechanical properties, potentially affecting biosensors performance.
For example, Tian et al. developed an optically modulated ionically conductive hydrogel to enable stimuli-responsive bioelectronics by combining Iron(II, III) oxide (Fe₃O₄) nanoparticles, polyacrylamide networks, and azo-benzene functionalized imidazole (AZIM) salt assemblies (Figure 5j).[71] The hydrogel's ionic conductivity was controlled through near-infrared light (808 nm), which generated localized heat via Fe₃O₄ nanoparticles, triggering the reversible disassembly of AZIM aggregates. This process provided a patterning resolution of 4 mm for electrodes.
Electrically, the hydrogel exhibited impedance modulation upon optical stimulation, with a conductivity increase proportional to light intensity and duration. Mechanically, the hydrogel demonstrated high stretchability and durability, maintaining performance under bending and repeated optical stimulation cycles, making it ideal for soft, bioinspired electronic applications.
Electronically Conductive Hydrogels
Electronically conductive hydrogels integrate the mechanical properties of soft, water-rich hydrogels with the electronic conductive materials.[73] Unlike ionically conductive hydrogels, which rely on ion movement, electronically conductive hydrogels incorporate electron-conducting materials—such as conductive polymers, carbon-based nanomaterials, or metallic nanomaterials—into their polymer network, enabling electron transport (Figure 5k).[74] This dual property makes electronically conductive hydrogels highly attractive for bioelectronics, particularly in implantable biosensors where both biocompatibility and electrical performance are critical. The conductivity of electrically conductive hydrogels results from the integration of conductive fillers within the hydrogel matrix. These fillers form interconnected pathways for electron transport, allowing the hydrogel to support electrical conduction, similar to traditional conductive materials. The hydrogel provides a flexible, biocompatible scaffold that can swell and deform like natural tissues, while the conductive fillers enable efficient electron flow.[75] Notable, when conductive polymer like PEDOT:PSS, PPy, or PANI are incorporated into hydrogel, the resulting materials combine the softness of hydrogel with the electronic conductivity of conductive polymers, offering a versatile platform for biosensor applications.
For instance, Lim et al. demonstrated tissue-like soft hydrogel with high mass-permeability and low impedance by incorporating PEDOT:PSS into poly(acrylamide) (PAAm) hydrogel.[6] The freeze-solidified PEDOT:PSS was mixed into the PAAm precursor and cross-linked together. PEDOT:PSS particles could be distributed inside the polymeric matrix to form PEDOT:PSS/PAAm hydrogel composite. In the following work, Shin et al. reported the homogeneously conductive PETOD:PSS/PANI/PAAm hydrogel composite with an exceptionally low impedance (≈21 Ω) and high conductivity (≈24 S cm−1).[76] The conductive hydrogel composite exhibited reliable adhesion on PU-based stretchable substrate via hydrogen bonding (Figure 5l).
Conductive Hydrogel Nanocomposites
Building upon ionically and electronically conductive hydrogels, significant advancements have been made by incorporating conductive nanomaterials into the hydrogel matrix, resulting in conductive hydrogel nanocomposites.[77] These nanocomposites leverage the unique properties of nanomaterials—such as exceptional electrical conductivity, mechanical strength, and high surface area—to enhance the overall performance of the hydrogel. By embedding conductive nanomaterials, such as carbon-base nanomaterials or metallic nanomaterials into hydrogels, these composites offer enhanced conductivity, mechanical softness, and functionality for demanding applications, including implantable biosensors.[78] The integration of nanomaterials addresses some inherent limitations of conventional conductive hydrogels, such as low conductivity or instability under physiological conditions.[79]
One of the primary advantages of conductive hydrogel nanocomposites is their ability to balance electrical and mechanical properties. The embedded nanomaterials create robust conductive networks, while the hydrogel matrix provides flexibility and biocompatibility. This synergistic combination allows for seamless integration with biological tissues, reducing mechanical mismatch and minimizing inflammation.[80] Furthermore, incorporating nanomaterials reinforces the hydrogel structure, enhancing long-term stability and mitigating degradation under physiological conditions.
For instance, Song et al. fabricated a highly stretchable and conductive hydrogel nanocomposite by integrating silver nanowire aerogels with poly(N-isopropylacrylamide) through a vacuum-assisted in situ polymerization process.[81] The silver nanowire aerogels formed a hierarchical ternary network that served as both a conductive framework and a large-scale crosslinker for the hydrogel. This design achieved exceptional electrical conductivity, reaching 93 S cm−1, with stable resistance even under 800% strain. Mechanically, the hydrogel exhibited remarkable tensile strength of 0.60 MPa at 1230% elongation (Figure 5n). Additionally, the composite demonstrated fast and efficient self-healing capabilities, achieving a 93% healing efficiency under near-infrared laser irradiation, ensuring durability for soft bioelectronics.
Similarly, Li et al. developed a multifunctional hydrogel nanocomposite by incorporating MXene nanosheets (Ti₃C₂Tx) into a polyacrylic acid (PAA) and amorphous calcium carbonate matrix (Figure 5o).[82] The hydrogel was fabricated using a biomineralization-inspired process where MXene nanosheets were dispersed within a hybrid polymer network. This structure exhibited high stretchability, achieving elongation of up to 450%, with a tunable Young's modulus of ≈300 kPa. Electrically, the hydrogel demonstrated low resistivity and exceptional sensitivity, with a gauge factor of 10.79 under strains up to 450%. Moreover, the material's self-healing capability restored electrical conductivity within 0.2 s, making it particularly suitable for wearable epidermal sensors and bioelectronics.
Recently, Lim et al. introduced a hydrogel nanocomposite featuring whiskered gold nanosheets (W-AuNSs) to achieve high conductivity and stretchability for soft bioelectronics.[83] Using a sequential formation method, a dry W-AuNS network was first fabricated, thermally treated, and then integrated into hydrogel matrices through curing (Figure 5p). This process established a stable percolation network of gold nanosheets, resulting in conductivity of up to 3304 S cm−1 and stretchability of 300%. The composite also maintained electrical stability over 100 cycles of 40% strain, making it highly suitable for bioelectronic applications such as in vivo tissue interfaces, where both durability and high performance are critical.
5 Conductive Fillers for Intrinsically Soft Conductors
The integration of conductive materials into soft, stretchable matrices is essential for ensuring both the functionality and mechanical compatibility of devices with biological tissues. Intrinsically soft conductive fillers are materials that combine electrical conductivity with flexibility, stretchability, and biocompatibility.[84] When embedded within a soft matrix, such as a polymer or hydrogel, these fillers create composites that maintain electrical conductivity even under deformation.[85, 86] This section examines various types of intrinsically soft conductive fillers, their mechanisms of action, and the advantages and challenges associated with their use.
The primary goal of incorporating conductive fillers into soft matrices is to achieve a balance between mechanical compliance and electrical performance. These fillers must retain their conductive properties under mechanical strain, a critical requirement for the reliable operation of biosensors exposed to continuous movement and deformation in the body.[87] Common materials used as soft conductive fillers include conductive polymers, carbon-based nanomaterials, metallic nanomaterials, and liquid metals. This section also discusses the fundamental principles, such as percolation network theory, to provide a comprehensive understanding of how these fillers function within composite systems.
5.1 Introduction to Percolation Network Theory
Percolation network theory is a fundamental concept that describes the transition from an insulating to a conductive state in composite material as the concentration of conductive fillers increases.[88] This theory is particularly relevant in the context of soft conductive composites, where conductive fillers are dispersed within an insulating matrix.
5.1.1 Percolation Threshold
The percolation threshold refers to the critical concentration of conductive fillers at which a continuous network of conductive pathways forms within the composite material.[89, 90] Below this threshold, the composite remains insulating because the conductive fillers are isolated and unable to facilitate electron flow. Once the percolation threshold is reached, the fillers establish an interconnected network spanning the entire material, enabling the flow of electric current, which causes a rapid increase in electrical conductivity.[91]
While exceeding the percolation threshold significantly enhances conductivity, excessive filler concentrations can adversely affect the composite's mechanical flexibility and softness.[92] Thus, precise control of filler concentration is critical to achieving a balance between electrical and mechanical properties. Achieving a low percolation threshold—where minimal filler is required to form a conductive network—is particularly desirable as it helps preserve the inherent flexibility and softness of the composite material.[93] Understanding and optimizing the percolation threshold is fundamental for the development of soft conductive composites with superior performance.[94]
5.1.2 Factors Influencing Percolation
Several factors influence the percolation threshold and the conductivity of the resulting composite, which must be carefully optimized to achieve desired performance. On of the most influential factors is the geometry of the conductive fillers. The shape and aspect ratio of fillers significantly affect their ability to form percolating networks. The high aspect ratio of the fillers increases the likelihood of creating interconnected pathways, thus reducing the percolation threshold and improving conductivity (Figure 6b,c).[95] The properties of the matrix also play a significant role. Characteristics such as viscosity, polarity, and filler-matrix interactions influence the dispersion and distribution of fillers, directly affecting the percolation threshold. A matrix that promotes strong filler-matrix interactions enhances filler distribution, which enables the formation of conductive pathways at lower filler concentrations. Filler concentration is another crucial factor.[96] While increasing the concentration of conductive fillers beyond the percolation threshold further enhances conductivity, excessive filler loading can compromise the mechanical properties of the composite, increasing stiffness and reduced flexibility. This trade-off highlights the need for balancing of conductivity and mechanical performance.
These factors collectively determine the formation of the percolation network, and their optimization is crucial for designing composite materials with both high conductivity and flexibility. For dynamic environments—such as wearable electronic devices—soft composites must maintain conductivity while withstanding repeated deformation. Therefore, selecting and arranging fillers is essential for achieving robust and reliable performance.
5.2 Conductive Carbon-Based Nanofillers
Carbon-based nanomaterials, CNTs, graphene, graphene oxides (GOs), graphite, and carbon fiber, are widely utilized as conductive fillers in soft composites due to their exceptional electrical, mechanical, and thermal properties. These materials have been widely studied for their potential to create highly conductive and flexible composites, particularly for biosensors and other electronic devices. This section will further explore these materials, providing examples of their applications and examining their role in advancing the performance and functionality of soft electronic and biosensing technologies.
5.2.1 1D Carbon Nanofillers: CNTs
1D carbon nanofillers, such as CNTs, possess elongated structures with high aspect ratios. CNTs are cylindrical structures composed of rolled-up sheets of graphene, with diameters ranging from a few to tens of nanometers (Figure 6d).[97] Renowned for their exceptional electrical conductivity, high aspect ratios, and mechanical strength, CNTs are effective conductive fillers in composite materials.[98] Their unique properties make them valuable in flexible and stretchable electronics, sensors, and energy storage devices.[99]
However, integrating CNTs into composites poses challenges due to their tendency to agglomerate caused by van der Waals interactions. This aggregation can lead to non-uniform conductivity and mechanical properties. To address this, techniques such as ultrasonication, chemical functionalization, and the use of surfactants, are employed to achieve homogeneous distribution of CNTs. For instance, Kim et al. developed a simple and cost-effective CNTs/PDMS nanocomposite for intrinsically soft electronics.[100] The fabrication process involved ultrasonication of CNTs in isopropyl alcohol to ensure uniform dispersion, followed by blending with PDMS and curing at 80 °C for 2 h (Figure 6e,f). This approach produced a well-dispersed CNTs network within the PDMS matrix, resulting in excellent electrical conductivity (sheet resistance of 2.03 Ω sq−1 with 8 wt.% CNTs) and stretchability up to 45% strain. The composite also demonstrated outstanding durability, maintaining electrical stability over 10000 strain cycles, making it suitable for applications in flexible circuits, strain sensors, and biopotential electrodes.
In another example, Kazemi et al. fabricated a stretchable thermoplastic vulcanizate nanocomposite using multi-walled CNTs as fillers in a maleic anhydride-grafted polyethylene matrix containing pre-vulcanized rubber particles (Figure 6g).[101] This segregated structural design achieved a low percolation threshold of 0.5 wt.% CNTs, enabling excellent electrical conductivity of up to 10−2 S cm−1 along with high mechanical stretchability. The composite withstood 1000 cycles of 10% strain while retaining 84.5% of its initial tensile strength. Furthermore, the CNTs network exhibited low resistance under stretching, with only a threefold increase in resistance observed at 50% strain. This combination of high conductivity, durability, and elasticity highlights the suitability of CNTs-based composites for flexible and soft electronics.
5.2.2 2D Carbon Nanofillers: Graphene
2D carbon nanofillers, such as graphene, GO, and reduced graphene oxide (rGO), are characterized by their thin, planar structures, and large surface areas. These properties make them ideal for forming efficient conductive networks while maintaining the mechanical flexibility of the composite. The high surface area of 2D fillers also enhances their interaction with the surrounding matrix, further contributing to improved conductivity and mechanical performance.
Graphene, a single layer of carbon atoms arranged in a two-dimensional honeycomb lattice, is widely recognized for its extraordinary electrical conductivity, mechanical strength, and flexibility.[102] Its high aspect ratio and large surface area make it an excellent conductive filler in soft composites, enabling applications in flexible electronics, sensors, and energy storage devices.[103, 104]
Despite its advantages, achieving uniform dispersion of graphene within a matrix remains challenging due to its tendency to agglomerate. Chemical functionalization can improve dispersion and compatibility with various matrices but may affect graphene's intrinsic properties. Another significant limitation is the scalable and cost-effective production of high-quality graphene, as techniques like chemical vapor deposition (CVD) and mechanical exfoliation often have limited throughput (Figure 6h).[105] Furthermore, graphene's biocompatibility and long-term stability in physiological environments require further investigation to ensure its suitability for biomedical applications.
GO, a derivative of graphene, features oxygen functional groups (e.g., hydroxyl, epoxide, and carboxyl) on its surface. These functional groups make GO hydrophilic and easier to disperse in aqueous solutions, improving its processability. However, the presence of oxygen groups reduces its electrical conductivity. rGO, produced by partially removing these oxygen groups, restores some conductivity while retaining the improved dispersibility of GO. GO and rGO have found applications in flexible electronics, sensors, and tissue engineering. For instance, rGO-based composites are used in glucose sensors, where their high surface area and moderate conductivity enhance performance. For instance, Liu et al. developed a highly stretchable graphene oxide/polyacrylamide nanocomposite hydrogel by utilizing GO nanosheets as cross-linkers through in situ free radical polymerization (Figures 6i).[106] This approach created a robust organic-inorganic network stabilized by hydrogen and ionic bonds. The hydrogel achieved a tensile strength of 385 kPa and an elongation at break exceeding 3400%, significantly outperforming conventional PAM hydrogels (Figure 6j). The well-dispersed GO nanosheets provided enhanced flexibility and toughness, enabling over 10× higher ductility. This PGH material demonstrated resilience under repeated stress, making it ideal for applications in soft robotics, bioelectronics, and flexible wearable devices. Recently, graphene fabrication technique using simple laser irradiation has been proposed, called laser-induced graphene (LIG). LIG is a highly conductive and porous material created by laser carbonization of a polymer substrate like PI. It offers excellent conductivity, lightweight properties, and scalability, making it suitable for flexible and wearable electronics. For example, Lu et al. used a cryogenic transfer method to integrate LIG with an ultrathin polyvinyl alcohol-phytic acid-honey hydrogel film to create stretchable and conductive nanocomposites (Figure 6k).[107] The cryogenic process at −196 °C enhanced the interfacial bonding between LIG and the hydrogel, allowing deflected cracks instead of straight ones during stretching. This design achieved a stretchability increase from 20% to over 110%, with a conductivity of 62 Ω sq−1 after transfer. The nanocomposites maintained mechanical integrity under repeated strain, showing potential for applications in soft bioelectronics such as wearable sensors and implantable cardiac patches.
5.2.3 3D Carbon Nanofillers: Graphite
3D carbon nanofillers, including graphite and CNT pillars, form complex, multidimensional networks that provide numerous contact points for electron transport. These 3D structures establish conductive pathways in multiple directions, maximizing the electrical performance of the composite. As a result, 3D fillers are particularly suitable for applications requiring high conductivity and robust structural integrity, such as in advanced biosensors or energy storage devices.
Graphite, a crystalline form of carbon composed of stacked layers of graphene sheets, offers several advantages as a primary filler for soft bioelectronics (Figure 6l).[108] These include low cost, scalability, and high electrical conductivity. Its layered structure facilitates efficient electron transport, making it suitable for applications requiring stable and reliable conductive pathways.[109, 110] Compared to CNTs and graphene, graphite is more readily available and easier to process. However, its bulkier structure and lower aspect ratio limit its flexibility and percolation efficiency at low concentrations. Unlike graphene's exceptional mechanical strength and CNTs’ high aspect ratio, graphite is less effective in forming stretchable composites. Additionally, achieving uniform dispersion of graphite within soft matrices remains a challenge. As a result, graphite is more commonly utilized as a precursor for graphene or in energy storage applications, rather than as a nanocomposite filler for flexible electronics.
Despite these limitations, innovative approaches have been developed to enhance the performance of graphite-based composites. Wang et al. synthesized graphite nanoplate (GN) nanocomposites using a surfactant-assisted ultrasonic exfoliation process, employing sodium carboxymethyl cellulose (CMC) to control stacking orientation (Figures 6m,n).[111] These nanocomposites were integrated into a PI substrate via a gap-coating method, forming GN/PI bi-layer films for electro-thermally driven actuators. The resulting GN films exhibited low resistivity (13.4–38.2 mΩ·cm) and high tensile strength (up to 20.3 MPa) with 9.1 wt.% CMC. This optimized composition enabled precise bending with a curvature of 2.2 cm⁻¹ and maximum bending angles of 270° under low voltages (3–48 V). The actuators maintained stability over 100 actuation cycles, demonstrating their potential for use in soft robotics and flexible electronics.
5.3 Conductive Metallic Nanofillers
Metallic nanomaterials, including nanoparticles, nanowires, and nanosheets, are widely utilized as conductive fillers in soft composites due to their exceptional electrical conductivity and tunable optical and electronic properties.[112, 113] These materials are particularly valuable for applications requiring both high conductivity and mechanical flexibility. The dimensional structure of metallic nanomaterials plays a critical role in their dispersion within composites and the formation of efficient percolation networks. Each structural form—0D, 1D, or 2D—offers unique properties and application potential, making the selection of appropriate material essential for optimizing the electrical and mechanical performance of composites.
5.3.1 0D Metallic Nanomaterials: Metal Nanoparticles
0D metallic nanomaterials, including silver (AgNPs), gold (AuNPs), copper (CuNPs), and platinum (PtNPs), are critical components in the development of soft implantable bioelectronics due to their unique electrical, chemical, and biological properties.[114] These zero-dimensional nanomaterials offer high electrical conductivity and biocompatibility, enabling applications in biosensors, neural interfaces, and energy storage for implantable devices (Figure 7a). Their small size facilitates incorporation into soft matrices, allowing for flexible and stretchable composites that integrate seamlessly with biological tissues.
AgNPs
AgNPs are extensively used in implantable bioelectronics due to their exceptional electrical conductivity and cost-effectiveness (Figure 7b).[115] Their high surface area-to-volume ratio enhances interaction with surrounding matrices, making them ideal secondary fillers to stabilize networks formed by primary conductive fillers. AgNPs are commonly synthesized through chemical reduction of silver salts, offering scalability and tunable nanoparticle size and shape. Eco-friendly green synthesis methods, using biological agents such as plant extracts or polysaccharides, are particularly attractive for biomedical applications due to reduced toxicity. However, AgNPs face challenges such as oxidative instability, which can degrade their electrical properties, and potential cytotoxicity, necessitating further research on stabilization techniques and biocompatibility.
AuNPs
AuNPs are highly valued in soft bioelectronics for their excellent electrical conductivity, chemical stability, and biocompatibility (Figure 7c).[116] Unlike AgNPs, AuNPs are resistant to oxidation, ensuring stable performance in long-term implantable applications. AuNPs also possess unique optical properties due to localized surface plasmon resonance, enabling their use in biosensing applications such as molecular imaging and cancer diagnostics. Their surface functionalization versatility allows AuNPs to bind with biomolecules or drugs, making them suitable for targeted drug delivery and biosensors. Common synthesis methods include citrate reduction for spherical particles and seed-mediated growth for complex shapes, allowing precise control over size and morphology to meet specific biomedical requirements.
CuNPs
CuNPs present an economical alternative to AgNPs and AuNPs, offering comparable electrical conductivity at a lower cost.[117] They also demonstrate notable thermal conductivity, broadening their applications in implantable energy devices and bioelectronics requiring efficient heat dissipation. Despite their advantages, CuNPs are highly prone to oxidation, which compromises their electrical performance. Protective coatings or functionalization with stabilizing agents, such as graphene or polymers, are employed to mitigate this issue. Green synthesis methods for CuNPs, similar to those used for AgNPs, are increasingly explored for their environmental benefits and potential in transient electronics. CuNPs have also been incorporated into bioresorbable composites for temporary implants, where their biodegradability is advantageous.
PtNPs
PtNPs are distinguished by their exceptional catalytic activity, chemical stability, and biocompatibility, making them ideal for specialized applications such as biosensors and energy-conversion devices.[118] In implantable bioelectronics, PtNPs are often used as electrode materials in fuel cells or enzymatic sensors, where their catalytic efficiency enhances performance. PtNPs are highly resistant to corrosion and degradation, ensuring long-term stability in physiological environments. Their biocompatibility has been leveraged in neural implants and drug delivery systems. However, the high cost of platinum limits its widespread use, confining it to applications requiring unmatched performance. Synthesis techniques such as chemical reduction and polyol reduction allow precise control over PtNP size and morphology, optimizing catalytic properties for specific applications.
5.3.2 1D Metallic Nanomaterials: Metal Nanowires
1D metallic nanomaterials, such as silver nanowires (AgNWs), gold nanowires (AuNWs), copper nanowires (CuNWs), possess high aspect ratios that enable the formation of efficient percolation networks at relatively low concentrations (Figure 7e).[119] Their elongated geometry allows them to form efficient percolation networks even at low concentrations, ensuring excellent electrical conductivity without compromising the mechanical flexibility of the composite. This ability to maintain electrical connectivity under mechanical strain makes metallic nanowires particularly valuable for implantable devices that experience continuous deformation. Embedded within soft matrices, nanowires can slide and rearrange, preserving conductive pathways even during stretching or bending, which is critical for long-term functionality in biomedical environments.
AgNWs
AgNWs are among the most extensively studied one-dimensional metallic fillers due to their inherent high conductivity and excellent aspect ratios (Figure 7f).[120] AgNWs can form efficient percolation networks at relatively low filler concentrations, resulting in composites with low percolation thresholds and high stretchability.[121] This makes them ideal for applications in implantable bioelectronics, such as flexible neural interfaces, stretchable electrodes, and biosensors. AgNWs are typically synthesized using methods like the polyol process, where silver precursors are reduced in a polyol solvent such as ethylene glycol, with polyvinylpyrrolidone (PVP) as a capping agent to promote anisotropic growth.[122] Reaction conditions, such as PVP concentration and temperature, allow precise control over nanowire length and diameter, optimizing them for specific biomedical applications.[123] Despite their advantages, AgNWs face challenges related to oxidation, which can degrade their conductivity over time. Protective coatings or surface modifications are often employed to enhance their stability, ensuring their suitability for long-term implantation.
AuNWs
AuNWs share the excellent conductivity of silver nanowires but offer superior chemical stability and biocompatibility, making them particularly suitable for implantable bioelectronics (Figure 7g).[124] Unlike silver, gold is resistant to oxidation, ensuring consistent performance over extended periods in physiological environments. AuNWs are used in applications such as stretchable biosensors, flexible circuits, and implantable neural interfaces, where stability and biocompatibility are critical. However, synthesizing high-aspect-ratio AuNWs remains a challenge due to gold's high surface diffusion and lack of a natural protective oxide layer. To overcome this, hybrid nanowires such as Au-AgNWs have been developed, where a silver core is coated with a gold sheath, combining the high conductivity of silver with the chemical stability of gold.[125] This approach has proven effective in creating robust one-dimensional networks for implantable applications.
CuNWs
CuNWs offer a low-cost alternative to silver and gold nanowires while maintaining excellent electrical and thermal conductivity (Figure 7h).[126] Their affordability and ease of synthesis make CuNWs attractive for cost-sensitive implantable devices, such as transient biosensors and energy storage components. However, like other copper-based nanomaterials, CuNWs are highly prone to oxidation, which significantly reduces their electrical performance over time. To address this limitation, CuNWs are often encapsulated in protective polymers or coated with stabilizing layers, such as graphene or silver, to improve their stability in physiological environments. CuNWs are typically synthesized through the polyol process or thermal decomposition, offering control over nanowire dimensions by adjusting reaction parameters. These methods enable the production of CuNWs with properties optimized for applications requiring stretchability and high conductivity, such as flexible bioelectronic implants.
5.3.3 2D Metallic Nanomaterials: Metal Nanosheets
2D metallic nanomaterials, such as silver nanosheets (AgNSs) and flakes, gold nanosheets (AuNSs), offer unique advantages for soft implantable bioelectronics, particularly in applications requiring flat, conductive networks or films (Figure 7i).[127] These materials feature broad, plate-like structures that provide a large surface area, promoting efficient charge and heat transfer to surrounding matrices. This property makes them particularly suitable for forming two-dimensional conductive networks in soft composites. However, the integration of 2D metallic nanomaterials into elastomeric matrices presents challenges, including their tendency to aggregate and their potential to disrupt the mechanical integrity of the matrix due to their broad, flat structures.
Several synthesis methods have been developed to optimize the properties of metallic nanosheets for bioelectronic applications. Liquid-phase exfoliation is widely used to produce thin nanosheets from bulk metals through the application of mechanical forces such as sonication. This method enables the production of large quantities of nanosheets with uniform thickness, and by selecting appropriate solvents and surfactants, aggregation can be minimized, improving their dispersion within elastomeric matrices. CVD is another popular approach, offering precise control over nanosheet thickness and lateral dimensions. This technique is ideal for applications requiring high-quality, defect-free nanosheets, as it enables layer-by-layer growth of ultra-thin metallic sheets. Wet chemical synthesis, involving the reduction of metal salts in solution with surfactants or capping agents, further allows for controlled growth of nanosheets with tailored morphologies and surface chemistries.
AgNSs and Silver Flakes
AgNSs (or silver flakes) are among the most widely studied 2D metallic nanomaterials for soft implantable devices (Figure 7j).[128] Their large lateral dimensions enable the formation of continuous, highly conductive pathways within polymeric matrices, providing excellent electrical performance even at low filler concentrations. AgNSs are used in neural interfaces and biosensors, where their high conductivity facilitates reliable signal transmission. Moreover, their ability to slide and maintain contact under mechanical deformation ensures stable performance in stretchable and flexible devices. However, challenges such as oxidation and high cost limit their broader applicability, necessitating protective coatings or functionalization to enhance stability in physiological environments.
AuNSs
AuNSs have emerged as promising materials for soft implantable bioelectronics due to their superior electrical conductivity, chemical stability, and biocompatibility (Figure 7k,l).[129, 130] Their large surface area promotes efficient integration with soft matrices, enabling robust percolation networks that maintain conductivity under mechanical strain. AuNSs are particularly suited for applications such as flexible electrodes, biosensors, and neural interfaces, where long-term stability and biocompatibility are critical. Despite their advantages, the high cost of gold remains a limitation, confining their use to high-performance devices where their unique properties justify the expense.[131]
5.4 Liquid Metals
Liquid metals are a class of materials that remain in a liquid state at or near room temperature. They exhibit high electrical conductivity, low viscosity, and excellent deformability, making them ideal for use in flexible and stretchable electronics. Eutectic gallium-indium (EGaIn) is an alloy of gallium and indium that remains in a liquid state at room temperature (melting point 15.5 °C). Galinstan is an alloy composed of gallium, indium, and tin that remains in a liquid state at room temperature and lower (melting point −19 °C). Both liquid metals exhibit high electrical conductivity and excellent deformability, making them suitable for use in flexible and stretchable electronics. Additionally, EGaIn and Galinstan are also non-toxic and biocompatible, making them ideal for biomedical applications.
For instance, Lee et al. developed intrinsically soft bioelectronics using EGaIn-based liquid metal microgranular particles via a meniscus-guided printing (MGP) method.[132] This process utilized polyelectrolyte-coated EGaIn particles suspended in acetic acid to ensure stability and conductivity (Figure 8a). The MGP method enabled direct, high-resolution patterning with a line width as fine as 50 µm on various substrates, achieving a conductivity of 1.5×106 S m−1 without post-processing. Mechanically, the printed electrodes demonstrated stretchability up to 500% strain and retained electrical performance after 10 000 cycles of 100% strain. This versatile approach facilitated the creation of soft, deformable devices, including wearable ECG sensors and e-skins (Figure 8b).
EGaIn nanocomposite have also been widely applied in soft implantable sensors. For instance, Nam et al. presented the development of a needle-like microfiber for intrinsically soft bioelectronics using EGaIn as the liquid metal core, encapsulated within a nanocomposite shell.[133] The microfiber achieves temporary stiffness via freeze-spraying, enabling minimally invasive implantation. The EGaIn core offers high conductivity (3.4×104 S cm−1) and strain-insensitivity, while the shell provides enhanced stretchability up to 800% and multifunctionality, including pH sensing and electrical stimulation (Figure 8c). The microfiber maintains performance across 10 000 cycles of 100% strain, making it suitable for applications in cardiac monitoring, gastric sensing, and muscular stimulation, ensuring biocompatibility and mechanical durability. In another example, Choi et al. developed a strain-adaptive fiber-interlocked electronic patch for intrinsically soft implantable bioelectronics using EGaIn as a key material (Figure 8d).[134] EGaIn was embedded in a self-healing polymer matrix to form a nanocomposite. This composite demonstrated high conductivity (1.8×104 S m−1) and exceptional stretchability, maintaining electrical performance under 500 stretching cycles at 50% strain. Additionally, the patch utilized an adhesive hydrogel layer for conformable tissue adhesion within 0.5 s, achieving a shear adhesion strength of 7.2 kPa. The patch effectively recorded cardiac signals in vivo for over four weeks without tissue damage, showing its potential for advanced cardiac monitoring applications.
Galinstan also exhibited as the promising candidate of intrinsically stretchable electronics, as the EGaIn. For instance, Mao et al. demonstrated galinstan embedded in Ecoflex elastomer for intrinsically soft electronics.[135] Galinstan offers high electrical conductivity (2.89 × 10⁻⁷ Ω m) and stretchability, integrated within channels measuring 0.5 mm thick and 1 mm wide. The system achieves bending angles of 70° under 3 A current and sustained durability over 2 million cycles. Low driving voltage (<1 V) ensures biocompatibility, while its elasticity enhances adaptability in implantable applications. This approach balances mechanical compliance and electrical functionality, enabling soft robotics and bioelectronic devices with superior safety and adaptability (Figure 8e). In other work, Gao et al. introduced galinstan-based intrinsically soft implantable electronics designed as biomolecular sensors.[136] Galinstan's high conductivity and flexibility were utilized in microfluidic channels embedded within PDMS, providing biocompatibility and mechanical durability. The device demonstrated excellent electrical properties, with a detection sensitivity of 0.0835 kPa⁻¹ and linear response up to 0.8 MPa. The system achieved high mechanical robustness, tolerating strains exceeding 200%, while maintaining operational stability under physiological conditions (Figure 8f). These characteristics were validated in animal models, illustrating the sensor's capability for precise biomolecular detection with minimal physiological disruption.
The patterning technique of the galinstan was also presented. For instance, Lin et al. focused on improving the wettability of galinstan, a gallium-based liquid metal, on metal substrates for its application in intrinsically soft electronics.[137] By treating copper surfaces with CuCl2 solution, the interfacial contact angle of galinstan reduced significantly from approximately 150° to nearly 0°, enabling spontaneous wetting (Figure 8g). This reactive-wetting behavior was attributed to the formation of Cu–Ga intermetallic compounds and metallic bonds, facilitated by surface treatment. The treated substrates demonstrated enhanced spreading and electrical conductivity, paving the way for flexible circuit fabrication (Figure 8h).
Despite their extraordinary high conductivity and low modulus, liquid metals face several challenges. One major issue is the formation of an oxide layer on their surface, which can reduce conductivity and hinder adhesion to other materials. Additionally, the integration of liquid metal requires care design and encapsulation strategies to prevent leakage into surrounding tissue due to its fluidity. Representative characteristics, advantages, and limitations of the conductive materials introduced above are described in Table 2.
| Material | Electric conductivity [S cm−1] | Stretchability [%] | Biocompatibility | Key advantages | Key limitations | |
|---|---|---|---|---|---|---|
| Conductive polymers | PEDOT:PSS | ≈100–500 | ≈50–100 | Moderate to high | High conductivity, tunable mechanical properties, stable in aqueous environments | Requires doping for optimal performance, lower conductivity than metals |
| PPy | ≈1–10 | ≈30–80 | High | Good biocompatibility, stable under physiological conditions | Poor mechanical robustness, prone to degradation over time | |
| PANI | ≈10-100 | ≈10-50 | Moderate | Simple synthesis, cost-effective | Susceptible to pH-induced conductivity changes | |
| Carbon nanomaterials | CNT | ≈1000-10000 | ≈100-300 | High | Ultra-high conductivity, excellent mechanical properties | Tendency to agglomerate, dispersion challenges |
| Graphene | ≈1000-5000 | ≈10-100 | High | High electrical conductivity, lightweight | Difficult large-scale processing, potential cytotoxicity | |
| Metallic nanomaterials | Silver | ≈10000-50000 | ≈200-500 | Low | Extremely high conductivity, well-established synthesis | Prone to oxidation, potential cytotoxicity |
| Gold | ≈4000-20000 | ≈100-300 | High | Biocompatible, oxidation-resistant | High cost, limited availability | |
| Copper | ≈5000–30000 | ≈100-400 | Moderate | High conductivity, low cost | Susceptible to oxidation, requires protective coating | |
| Liquid metals | EGaIn | ≈10000-30000 | Fluidic | High | Excellent stretchability, self-healing | Requires encapsulation, potential toxicity |
| Galinstan | ≈10000-30000 | Fluidic | High | High conductivity, room temperature fluidity | Oxidation-sensitive, difficult patterning | |
|
Conductive hydrogel |
Ionically conductive | ≈0.01-1 | ≈200-1000 | High | Biocompatible, excellent hydration properties | Limited conductivity, prone to dehydration |
| Electronically conductive | ≈10-100 | ≈100-500 | High | Combines electronic and ionic conductivity | Stability in physiological environments can be challenging | |
6 Fabrication Process of the Unconventional Materials
Soft and stretchable bioelectronics require fabrication techniques that accommodate the inherent deformability of intrinsically soft materials while preserving their electrical and mechanical integrity. Unlike conventional lithographic methods optimized for rigid substrates, the patterning of soft electronic materials necessitates approaches that minimize mechanical stress and prevent structural damage. Various unconventional patterning techniques, such as soft lithography, inkjet printing, 3D printing, and laser-based patterning, have been developed to enable precise and scalable fabrication of soft bioelectronic devices. This chapter explores these methods in detail, examining their advantages, limitations, and practical challenges related to resolution, scalability, and reproducibility for biomedical applications. Additionally, a comparative analysis of these techniques is included to assist researchers in selecting the most suitable method based on device requirements (Table 3).
| Technique | Resolution | Scalability | Material compatibility | Application suitability |
|---|---|---|---|---|
| Photolithography | High (sub-µm) | Low | Photoresists, stretchable polymers | Microelectronics, MEMS integration |
| Soft lithography | High (1–10 µm) | Moderate | PDMS, elastomers, biomaterials | Microfluidics, biointerfaces |
| Inkjet printing | Moderate (10–50 µm) | High | Conductive inks, polymers, hydrogels | Flexible circuits, biosensors |
| 3D printing | Low (50–200 µm) | Moderate | Hydrogels, elastomers, conductive polymers | Customized implants, tissue scaffolds |
| Laser patterning | Moderate–High (1–100 µm) | High | Polymers, thin films, graphene | Wearable electronics, rapid prototyping |
6.1 Photolithography
Microelectronic mechanical systems (MEMS) are miniaturized devices that integrate mechanical elements, sensors, actuators, and electronics onto a common silicon substrate, traditionally leveraging photolithography for precise patterning.[138, 139] Photolithography involves using light to transfer geometric patterns from a mask onto a photosensitive material, enabling the creation of micro- and nano-scale features with exceptional accuracy (Figure 9a). While MEMS technology has historically been associated with rigid materials like silicon, recent advancements have focused on adapting MEMS techniques for unconventional soft materials, such as elastomers, hydrogels, and polymers, to develop flexible and stretchable biosensors. For soft bioelectronics, modified photolithography techniques have been developed using elastomeric substrates such as PDMS and photocurable polymers. In particular, photocrosslinkable elastomers (e.g., polyurethane acrylates) enable direct photopatterning of soft, stretchable circuits. However, the brittle nature of conventional photoresists can limit their applicability to highly deformable substrates. Advances in soft photolithography, such as the use of mechanically adaptive resists, have improved compatibility with stretchable materials.
One of the primary limitations of photolithography in soft electronics is scalability. Unlike printing-based approaches, photolithography typically requires mask alignment and cleanroom environments, increasing fabrication costs and complexity. Recent developments in maskless lithography and digital projection-based methods have aimed to address these scalability concerns, enabling more flexible and cost-effective fabrication workflows (Figure 9b).[140]
6.2 Soft Lithography
Soft lithography encompasses a set of microfabrication techniques that utilize elastomeric stamps, molds, or masks to pattern soft materials. Unlike traditional photolithography, which is typically limited to rigid substrates like silicon, soft lithography is particularly effective for creating micro- and nano-scale features on flexible and stretchable materials, such as PDMS, hydrogels, and other elastomers.[141, 142] These methods are particularly advantageous for bioelectronics because they allow for conformal contact with flexible substrates, preserving structural integrity during fabrication. This versatility makes soft lithography a critical tool in fabricating components for soft bioelectronics and biosensing applications.
One widely used techniques in soft lithography is microcontact printing.[143] This method employs a patterned elastomeric stamp, typically made from PDMS, to transfer an “ink”—which may include biomolecules, nanoparticles, or polymers—onto a substrate in a precise pattern (Figure 9c). The patterned stamp is created by casting PDMS against a master mold that defines the desired design. Microcontact printing is particularly effective for fabricating functionalized surfaces, such as biosensor arrays, where the patterned deposition of bioactive molecules is critical (Figure 9d).[144] However, achieving reproducible feature sizes and ink transfer efficiency remains a challenge. Variations in stamp deformation and ink loading can introduce patterning inconsistencies, particularly for nanoscale features. To improve reproducibility, advanced ink formulations and controlled stamp surface modifications have been explored.
Another key technique is replica molding, which involves casting an elastomer like PDMS against a master mold to replicate the mold's surface features.[145] This method enables the production of micro- and nano-scale structures with high fidelity and is commonly used to fabricate components for microfluidic devices. These devices, including microchannels and microvalves, play a crucial role in biosensing applications, allowing for controlled fluid flow and the detection of biological markers in small sample volumes. While replica molding offers excellent resolution, scalability remains a challenge due to the time-intensive nature of mold fabrication and material curing. Additionally, mold degradation over repeated use can affect pattern fidelity. Emerging high-throughput replication techniques, such as roll-to-roll soft lithography, are being investigated to enhance scalability for large-area bioelectronic manufacturing.
Microtransfer molding represents a variant of soft lithography in which an elastomeric mold is filled with a liquid precursor of the desired material.[146] Once cured, the solid microstructures are transferred onto the substrate when mold is removed. This technique is particularly valuable for creating complex three-dimensional microstructures, such as microchannels and microelectrodes, which can be integrated into flexible substrates for wearable biosensors and lab-on-a-chip devices.
Despite its many advantages, soft lithography faces challenges, such as ensuring high-fidelity pattern transfer, maintaining the mechanical stability of soft materials during fabrication, and addressing deformation or swelling of the elastomeric molds. Additionally, scaling up the process for high-throughput commercial production remains difficult. To overcome these limitations, recent advancements in soft lithography have focused on developing more robust elastomeric materials, improving patterning precision, and integrating soft lithography with complementary fabrication methods. For instance, hybrid processes that combine soft lithography with inkjet printing or laser cutting have enhanced the scalability and precision of microfabrication, paving the way for broader applications in flexible and wearable bioelectronics.
6.3 Printing Technique
Printing technique, including inkjet printing and 3D printing, uses the conductive ink printed via nozzle as programmed. Inkjet printing deposits tiny droplets of material onto a substrate with precision, making it suitable for printing conductive inks, biological materials, and other functional substances onto soft substrates (Figure 9e).[147] This technique is commonly used to fabricate flexible electronics, biosensors, and microfluidic devices.[148, 149]
Inkjet printing involves the deposition of tiny droplets of functional ink onto a substrate in a digitally controlled manner. This method is particularly useful for fabricating soft bioelectronics because it enables low-temperature processing and direct patterning without the need for masks or etching steps. Conductive inks, including those based on silver nanoparticles, carbon nanotubes, and PEDOT:PSS, can be printed onto soft polymeric substrates to create stretchable and biocompatible circuits. However, a key challenge in inkjet printing is achieving uniform ink distribution and adhesion on soft, hydrophobic materials. Inconsistent droplet spreading can lead to defects in printed patterns, affecting device performance. To address this, surface treatments such as plasma activation or the use of surfactant-based ink formulations have been employed to improve wetting properties. Additionally, multi-pass printing strategies and sintering techniques are used to enhance conductivity and pattern resolution. For example, conductive inks can be printed onto flexible substrates to form stretchable circuits, while biological inks are used to deposit cells or biomolecules onto scaffolds for tissue engineering (Figure 9f).[150] Inkjet printing's high precision and control make it a versatile tool for creating complex and functional patterns in soft biosensors.
3D printing, or additive manufacturing, is a process that builds three-dimensional structures layer by layer, offering remarkable design flexibility for soft materials used in biosensors (Figure 9g).[151, 152] This technique enables the fabrication of complex, customized geometries that conform to the contours of the body, making it particularly advantageous for wearable and implantable devices (Figure 9h).[153] Despite its versatility, 3D printing of soft electronic materials faces limitations in resolution and printing speed. The viscosity of printable inks and the precision of extrusion nozzles significantly impact the fidelity of printed structures. Furthermore, achieving high electrical conductivity in printed conductive materials often requires post-processing treatments, such as chemical sintering or thermal annealing. These additional steps can introduce compatibility issues with soft, temperature-sensitive materials.
Extrusion-based 3D printing, such as fused deposition modeling, involves the controlled deposition of material through a nozzle to construct structures layer by layer.[154, 155] Commonly used with thermoplastic elastomers and hydrogels, this technique is widely employed in fabricating soft implants, wearable sensors, and tissue scaffolds. For instance, hydrogels can be printed into scaffolds that support cell growth for tissue engineering, while soft elastomers can be used to create flexible sensors that adapt to the body's shape.
Stereolithography (SLA) employs a UV laser to cure liquid photopolymers layer by layer, forming solid structures with exceptional detail and smooth surfaces.[156, 157] SLA is particularly suited for applications requiring fine features, such as implantable biosensors and microfluidic devices. For example, SLA can produce intricate microchannels in photopolymers, which are critical components of lab-on-a-chip devices for biosensing applications. Its high resolution and precision make it ideal for creating detailed features in soft bioelectronics.
6.4 Laser Cutting
Laser cutting is a versatile fabrication technique that uses focused laser beams to precisely cut or engrave materials.[158, 159] Laser cutting can pattern flexible materials such as hydrogels and elastomers with high precision, using a high-powered laser to cut the material into the desired shape or form a fine pattern on the surface. Since laser cutting is a non-contact method that does not physically apply pressure to the material, it is ideal for creating precise patterns in soft, flexible materials, especially at high resolution. It also benefits from the ability to cut fine details with precision and the flexibility to work with different thicknesses or types of material (Figure 9i, left and middle).[160] For soft materials used in biosensors, laser cutting offers a high degree of precision and control, enabling the fabrication of complex geometries and fine features that are difficult to achieve with mechanical cutting tools.
The main challenges in laser cutting and engraving of soft materials include controlling the depth and precision of the cuts, minimizing thermal damage to the surrounding material, and preventing deformation or warping of the soft substrates. Additionally, the choice of laser parameters, such as wavelength, pulse duration, and power, must be carefully optimized for each material. Recent innovations in laser technology, such as the development of ultrafast lasers and advanced beam-shaping techniques, have significantly improved the precision and versatility of laser cutting for soft materials. These advances have enabled the fabrication of more complex and intricate biosensors, with features that were previously difficult or impossible to achieve (Figure 9j).[161]
Recently, laser-induced graphene (LIG) has emerged as a powerful technique for fabricating conductive networks directly on soft substrates (Figure 9i, right).[162, 163] By using laser beams to carbonize specific regions of polymeric materials, LIG allows the creation of highly conductive graphene patterns. This method is particularly attractive for soft bioelectronics, as it enables the integration of flexible conductive pathways without additional processing steps, further enhancing device functionality.
7 Further Functionalization for Long-Term Biocompatibility
The long-term success of soft implantable biosensors hinges not only on their immediate biocompatibility but also on their ability to maintain functionality and stability over extended periods within the body. Chronic biocompatibility involves minimizing the body's immune response, preventing device degradation, and ensuring that the implant remains integrated with the surrounding tissues without causing inflammation or fibrosis. To achieve this, various functionalization strategies can be employed, including enhancing the softness of the device, modifying surface properties to increase hydrophilicity, applying anti-inflammatory coatings, and using biomimetic surface coatings. Beyond these considerations, the long-term mechanical durability of implantable biosensors is crucial, as repeated physiological movements and strain can degrade device performance over time. Factors such as fatigue resistance, encapsulation integrity, and resistance to delamination must be addressed to ensure stable operation under chronic conditions. This section explores these strategies in detail.
7.1 Softness Enhancement
Softness is a critical factor in the design of implantable biosensors, particularly when these devices are interfacing with soft tissues such as the brain, heart, or skin. Enhancing the softness of the device ensures that it can conform to the natural contours of the tissue, reducing mechanical mismatch and minimizing the risk of chronic irritation or inflammation. Choosing ultrasoft materials would be an option for maximizing softness of the materials. Especially, hydrogels, with their high water content and tissue-like softness, are ideal materials for enhancing the softness of implantable devices. Hydrogels can be engineered to have varying degrees of softness by adjusting their crosslinking density or by incorporating softening agents. These materials are particularly useful in applications where the device needs to mimic the mechanical properties of soft tissues.
For instance, Lim et al. developed a tissue-like, ultrasoft bioelectronics interface by using a low-impedance PAAm hydrogel functionalized with PEDOT:PSS.[6] The hydrogel achieved superior mass-permeability and conformability due to its ultrathin structure (150 µm), enabling close integration with human tissue. The functionalization reduced impedance significantly to below 1 kΩ at low frequencies, enhancing bio-signal fidelity. Additionally, the hydrogel's soft mechanical properties (Young's modulus of 8 kPa) matched that of tissue, minimizing inflammation, making seamless contact, and improving biocompatibility, while ensuring efficient bioanalytical applications such as transcutaneous oxygen monitoring and localized drug delivery (Figure 10a).
Despite these advancements, one of the critical challenges is ensuring that mechanically soft materials maintain their performance over long-term implantation. Hydrogels, for example, are prone to dehydration and degradation, which can alter their mechanical and electrical properties over time. A promising approach to addressing this issue is the incorporation of double-network hydrogels, which combine a rigid primary polymer network with a softer secondary network to enhance durability while maintaining compliance. Furthermore, hybrid hydrogel-elastomer systems have been developed to improve both softness and long-term mechanical resilience by integrating hydrogels with stretchable elastomer matrices.
In another example, Li et al. introduced a generalizable soft interlayer design to achieve tissue-level softness in intrinsically stretchable bioelectronics, enhancing flexibility and biocompatibility.[39] Utilizing SEBS as a soft interlayer between functional layers and an ultrasoft hydrogel substrate, the effective modulus of the devices was reduced to 5.2 kPa (Figure 10b). This significant reduction improved conformability and minimized foreign body responses. Mechanically, the device maintained structural integrity under 100% strain for 1000 cycles. In vivo tests demonstrated reduced scar tissue formation (18.3 µm thick) and stable electrophysiological performance, making the device highly suitable for long-term implantable bioelectronic applications.
For another strategy, ultrathin structure or microstructure of the substrate would enhance the biocompatibility of materials. The softness of a device can be enhanced by incorporating microstructure or porous designs that allow the material to deform more easily under mechanical stress. For instance, Xu et al. introduced the application of microstructured soft materials to enhance the flexibility and biocompatibility of intrinsically soft implantable bioelectronics.[164] By engineering the polymer network with a bottlebrush architecture and incorporating single-walled carbon nanotubes, the material achieves an ultralow Young's modulus (2.98–10.65 kPa), closely matching soft biological tissues (Figure 10c). The soft nanocomposite exhibited excellent conductivity (2.68–13.78 S m−1) while retaining mechanical resilience under strains exceeding 100%. Furthermore, the solvent-free elastomer demonstrates robust environmental stability and adhesive properties, making it suitable for extended use in implantable devices. In another example, Lee et al. developed ultrasoft implantable bioelectronics using nanomesh materials composed of polyurethane nanofibers and conductive gold layers, encapsulated with parylene.[165] The nanomesh achieved a low effective Young's modulus of 0.274 MPa, closely mimicking soft biological tissues, and maintained stable impedance (3.61 kΩ at 100 Hz) under 20% strain. This softness minimized mechanical interference, allowing cardiomyocytes to contract naturally with a maximum local strain of 5.21% (Figure 10d). Additionally, the devices provided reliable field potential monitoring for 96 hours without degradation, demonstrating enhanced biocompatibility and functionality for long-term in vivo applications in electrophysiological sensing.
However, while these materials have demonstrated significant improvements in mechanical compatibility, clinical translation remains a challenge. Many hydrogel-based devices exhibit swelling in physiological environments, which can alter their mechanical and electrical properties over time. Additionally, ensuring long-term adhesion to tissues without inducing fibrosis remains an area of active research. Further studies evaluating long-term stability in preclinical and clinical models will be essential for advancing these materials toward clinical use.
7.2 Chemically Stable Materials
Improving the chemical stability of implantable devices is critical for ensuring long-term performance and biocompatibility. One effective strategy involves using inherently stable materials, such as gold, platinum, iridium, and titanium, to construct electrodes. These materials exhibit excellent resistance to corrosion and oxidation, even in harsh physiological environments, making them ideal for prolonged use in biomedical applications.
For devices incorporating reactive materials like silver or copper, surface modification through coating with chemically stable materials can significantly enhance their stability. Coating techniques are commonly employed to apply layers of gold, platinum, or iridium onto the surfaces of reactive components. These coatings act as protective barriers, preventing oxidation and degradation while preserving the underlying material's electrical conductivity and mechanical properties. For instance, Haro et al. enhanced the biocompatibility of implantable bioelectronics by chemically stabilizing electrode surfaces with electroplated platinum coatings.[166] Using cyclic voltammetry in chloroplatinic acid baths, they achieved platinum film thicknesses up to 7 µm, with roughness ranging from 150 to 500 nm. The coatings reduced impedance to 388 Ω at 2 µm thickness, compared to 982 Ω for sputtered platinum, and improved corrosion resistance, lowering dissolution rates from 38.8 to 7.8 ng C−1 (Figure 10e). These improvements minimized tissue irritation and prolonged electrode lifespan under biphasic stimulation, making the device more suitable for long-term implantation. In addition to improving chemical stability, mechanical integrity under long-term implantation must also be addressed. Thin-film metal coatings are prone to cracking under cyclic strain, leading to performance degradation. To mitigate this, recent research has focused on the development of nanostructured metal composites and alloy-based coatings that enhance both chemical and mechanical stability. For example, Kim's group demonstrated the use of Ag-Au core-shell nanowires[125] and Ag–Au–Pt core–shell–shell nanowires[167] within an elastomeric matrix to create a stretchable, low-impedance, and biocompatible nanocomposite for bioelectronics applications (Figures 10f,g). The Pt outer shell is embossed to maximize charge transfer and reduce impedance (166.5 Ω at 1 kHz), and enhance conductivity (≈11000 S cm−1) and stretchability (≈500%). The Au–Pt coating effectively prevents Ag ion leaching, ensuring cytocompatibility.
Despite these advancements, long-term chemical stability under dynamic physiological conditions remains a concern. Encapsulation strategies that combine chemical stability with flexibility are needed to ensure that electrodes remain functional over years of implantation. Additionally, regulatory approval for new material coatings often requires extensive safety validation, including cytotoxicity, hemocompatibility, and in vivo biostability testing.
7.3 Anti-Inflammatory Material Coating
Chronic inflammation is a significant challenge for implantable devices, as it can lead to fibrosis, device encapsulation, and ultimately, failure of the implant. To mitigate this issue, devices can be coated with anti-inflammatory materials that actively suppress the body's immune response, promoting long-term biocompatibility.
7.3.1 Immunosuppressive Coatings for Controlled Drug Release
Applying immunosuppressive drug-releasing layers on the device surface is a widely explored strategy to mitigate inflammation and enhance biocompatibility. These coatings are designed to release anti-inflammatory agents, such as dexamethasone or rapamycin, at the implantation site, reducing immune responses that could lead to fibrotic encapsulation or device failure. Techniques like polymer encapsulation or nanoparticle-based systems are often used to control the release kinetics of these drugs, ensuring sustained suppression of inflammation over extended periods. For instance, Du et al. developed a novel drug-eluting stent coating featuring a hydrophobic core and hydrophilic shell nano/micro particle system to optimize drug release for vascular repair (Figure 10h).[168] Using coaxial electrospraying, the coating integrated docetaxel in a poly(lactic-co-glycolic acid)-based core and a platelet receptor antibody in a chitosan-based shell. This configuration achieved biphasic drug release: the antibody showed rapid early release, facilitating platelet inhibition, while docetaxel exhibited a slower release to suppress smooth muscle proliferation. The system demonstrated enhanced reendothelialization, reduced platelet activation, and lower neointimal hyperplasia in vivo.
However, achieving a sustained and controlled release remains challenging. Burst release of drugs can lead to suboptimal long-term effects, while insufficient release may fail to suppress inflammation effectively. Novel polymeric delivery systems, such as nanoparticle-based encapsulation or hydrogels with tunable degradation rates, offer promising approaches for improving controlled drug delivery for long-term implantation.
7.3.2 Antibiofouling Surfaces
Antibiofouling strategies aim to prevent the adhesion of proteins, cells, and microbes on the device surface, which can trigger inflammatory responses. Zwitterionic surfaces, composed of materials with both positive and negative charges, create a hydration layer that resists biofouling by repelling nonspecific protein adsorption and microbial attachment. For example, zwitterionic polymers such as poly(carboxybetaine) or poly(sulfobetaine) can be grafted onto device surfaces, forming highly biocompatible coatings that minimize immune activation. Specifically, Qin et al. utilized zwitterionic materials, specifically 2-methacryloyl phosphorylcholine (MPC), to coat implant surfaces for enhanced biocompatibility (Figure 10i).[169] By grafting a stable MPC gel layer onto PDMS surfaces via surface-initiated atom-transfer radical polymerization, the coating achieves superhydrophilicity with a water contact angle reduced from 110° to 15°. This modification suppresses protein adsorption, minimizes macrophage activation, and reduces bacterial adhesion by over 70% compared to uncoated PDMS. The mechanical properties, with elastic moduli spanning 0.5–42 kPa, align with soft tissue. This innovative approach offers a durable, antifouling surface ideal for minimizing inflammatory responses and improving implant longevity.
In addition, nanostructured surfaces, such as those with nanopillars, nanowires, or nanogrooves, can physically deter biofouling. These surfaces mimic natural antifouling structures, such as those found in lotus leaves or shark skin, creating topographies that disrupt microbial adhesion and protein accumulation. For implantable devices, combining zwitterionic coatings with nanostructured features provides a dual mechanism for reducing inflammation and enhancing biocompatibility. These strategies improve device performance by maintaining surface cleanliness and preventing biofilm formation, crucial for long-term functionality in vivo. For example, Zhang et al. focused on integrating mechano-bactericidal nanopillar arrays and multiplexed strain sensors into a smart polymer foil for orthopedic implants (Figure 10j).[170] The outer surface uses nanopillars mimicking cicada wings, optimized to a diameter of ≈100 nm, a height of ≈400 nm, and a pitch of ≈240 nm. These structures achieved over 99% bacterial clearance (E. coli, P. aeruginosa, and S. aureus) in vivo, while maintaining biocompatibility with mammalian cells. Meanwhile, single-crystalline silicon nanomembranes on the inner surface provide strain mapping with 0.01% sensitivity under low voltage, ensuring conformal application and stability over 8 weeks in vivo.
While these coatings have shown promise in preclinical studies, their long-term stability and mechanical robustness under physiological conditions remain critical concerns. Clinical translation will require extensive in vivo validation to ensure that these materials maintain their antifouling properties over extended implantation periods.
7.4 Biomimic Surface Coating
Biomimetic surface coatings are designed to mimic the natural properties of biological tissues, promoting integration and reducing the likelihood of an adverse immune response. These coatings can be engineered to replicate the extracellular matrix (ECM), present bioactive molecules, or create surfaces that interact favorably with cells.
Bioactive molecule coatings involve the immobilization of signaling proteins, growth factors, or peptides onto the surface of the device. These molecules interact with cells to promote specific biological responses, such as angiogenesis, wound healing, or immune modulation. Bioactive molecule coatings are used in implants that require active tissue integration or healing, such as vascular grafts, bone implants, and wound dressings. By presenting bioactive signals to the surrounding cells, these coatings help to guide the healing process and reduce the risk of fibrosis or chronic inflammation.
ECM-mimicking coatings are designed to replicate the composition and structure of the natural extracellular matrix, providing a familiar environment for cells. These coatings can include proteins like collagen, laminin, or fibronectin, as well as glycosaminoglycans and other ECM components. ECM-mimicking coatings are used in a variety of implants, including neural interfaces and tissue engineering scaffolds. By providing a surface that closely resembles the natural environment, these coatings help to promote cell adhesion, proliferation, and differentiation, leading to better integration and reduced immune response.
For instance, Kang et al. introduced an advanced approach for creating biomimetic soft implantable bioelectronics using collagen and fibrin-based hydrogels to enhance biocompatibility and wound healing (Figure 10k).[171] These materials mimic the ECM of natural skin, accelerating tissue regeneration while providing structural and mechanical properties similar to native tissues. The hydrogel exhibited a Young's modulus of ≈70.8 kPa, matching sciatic nerve tissue, and showed significant ion transport properties. Histological and immunofluorescence analyses revealed reduced fibrosis and inflammation when hydrogel-coated electrodes were implanted, highlighting minimal foreign body reactions and effective neural stimulation. In another example, Wang et al. focused on enhancing the biocompatibility and stability of intrinsically soft implantable bioelectronics by coating electrodes with rat red cell membranes (RCMs) supported by ionic liquid 1-butyl-2,3-dimethylimidazolium hexafluorophosphat films (Figure 10l).[172] The RCM layer reduced cortical tissue damage, evidenced by a 38% reduction in glial scarring compared to non-coated electrodes over 28 days. Electrochemical tests demonstrated the coated electrodes' superior charge transfer resistance (147.3 Ω post-implantation), indicating stable ion transport and minimized tissue inflammation. These findings highlight RCM-coated electrodes as a promising strategy for chronic bioelectronics with improved biocompatibility and electrical performance.
Recent advancements have explored the integration of biological tissues with electrode surfaces to enhance biocompatibility and functionality. Such as a study investigating the bioinspired coating of silicone implant surfaces using human adipose-derived stem cells (hASCs) and itaconic acid-conjugated PDMS (IA-PDMS) to enhance biocompatibility and minimize fibrotic responses (Figure 10m).[173] The hydrophilic IA-PDMS surface significantly improved hASC adhesion and proliferation compared to unmodified PDMS, as demonstrated by increased cell viability (p < 0.05). Implantation in rats over 60 days showed a marked reduction in capsule thickness and collagen density for hASC-coated surfaces, particularly hASC-IA-PDMS (p < 0.0001), along with lower pro-inflammatory cytokine release and greater M2 macrophage polarization.
Another innovative strategy is constructing electrodes directly using biological tissues. For instance, conductive hydrogels embedded with biological components or hybrid materials that incorporate living cells into the electrode structure have been investigated. For instance, in the study by Adewole et al. introduces “living electrodes” using long-projecting axonal tracts encapsulated in soft hydrogel microcolumns for brain-machine interfaces (Figure 10n).[174] These electrodes leverage microtissue-engineered neural networks (µTENNs), comprising living cortical neurons and axons, to improve biocompatibility, synaptic integration, and signal specificity. The fabrication involves agarose-based microcolumns seeded with neuronal aggregates and extracellular matrix, enabling controlled axonal growth at rates up to 1101.8 ± 81.1 µm d−1 in bidirectional constructs. In vivo, implanted µTENNs showed neuronal survival, axonal integrity, and synaptic connections with host tissue, alongside minimal inflammatory response. This novel strategy significantly enhances biocompatibility and stability over conventional neural interfaces.
8 Minimally Invasive Implantation and Fixation of Soft Bioelectronics
The success of soft implantable biosensors relies not only on their material composition and functionality but also on the techniques used for their implantation and fixation within the body. Minimally invasive implantation methods are crucial for reducing tissue damage, minimizing the immune response, and ensuring that the device maintains close contact with the target tissues or organs. Furthermore, the long-term fixation of these devices is essential to prevent movement, which could lead to signal degradation, mechanical failure, or adverse biological reactions. This section explores the various strategies and techniques used for the minimally invasive implantation and chronic fixation of soft bioelectronics.
8.1 Minimally Invasive Implantation Procedure
Minimally invasive implantation techniques aim to introduce soft bioelectronics into the body with minimal disruption to surrounding tissues. These methods are particularly important for applications where the device must be implanted in sensitive areas, such as the brain, heart, or peripheral nerves. The following subsections discuss different approaches to achieving minimally invasive implantation.
8.1.1 Implantation with the Rigid Shuttle
One of the most common techniques for implanting soft bioelectronics is using a rigid shuttle. This method involves temporarily mounting the flexible device onto a rigid structure that facilitates its insertion into the body. The rigid shuttle, often made from biocompatible materials such as stainless steel or silicon, acts as a carrier for the soft device. The shuttle is inserted into the target tissue, creating a path for the soft bioelectronics. Once the device is in place, the shuttle is removed, leaving the flexible device implanted. This technique is widely used for neural implants, such as flexible electrodes for brain-machine interfaces. The rigidity of the shuttle allows the delicate and flexible electrodes to be accurately positioned within the brain, where they can record neural activity or stimulate neurons.
For instance, Rho et al. presented a transient shuttle-based implantation strategy utilizing Poly(vinyl alcohol) (PVA)-coated mesh electrodes to achieve minimally invasive and chronic implantation of soft electronics in the brain.[175] The PVA shuttle initially provides rigidity (3.59 nN m2) for penetration and dissolves into flexibility (3.33 pN m2) post-insertion, reducing mechanical mismatch and immune response (Figures 11a,b). The fabrication involves encapsulating SU-8-based electrodes in PVA, which dissolves within 6 min in Phosphate buffered saline at 37 °C. This process minimizes insertion-induced stress and eliminates withdrawal damage. The system maintains stable electrode positioning, preventing tissue displacement and allowing long-term brain mapping.
8.1.2 Temporal Rigidity Changes Before/After Implantation
To overcome the challenges associated with using a rigid shuttle, researchers have developed materials and techniques that allow the device to change its rigidity temporally—being rigid during implantation and becoming soft and flexible afterward.[176] Shape-memory polymers and liquid metals are class of materials that can switch between a rigid and a flexible state in response to external stimuli, such as temperature. For example, a device made from SMPs and/or liquid metals can be temporarily stiffened at low temperatures to facilitate insertion, and once implanted, the body's natural heat causes the material to soften, conforming to the surrounding tissues. Another approach involves the use of hydrogels that swell and soften upon exposure to physiological fluids. These hydrogels are initially dehydrated and rigid, making them easier to implant. After implantation, the hydrogels absorb water from the body, expanding and becoming softer, which helps to reduce mechanical stress on the surrounding tissues.
Alternatively, Nam et al. recently showcased the use of a modulus-shifting material for minimally invasive implantation of intrinsically soft bioelectronics.[133] By utilizing a phase-convertible liquid metal core encapsulated in a nanocomposite shell, the microfiber transitions from a solidified state for tissue penetration (145 MPa elastic modulus) to a soft, tissue-compatible state post-implantation (13 MPa elastic modulus). Freeze-spraying facilitates stiffness control, enabling precise deployment without surgical tools (Figure 11c,d). The device maintains its conductivity (3.4 × 10⁴ S cm−1) and mechanical integrity during phase shifts, ensuring reliable implantation and seamless integration with dynamic tissues such as the heart and muscles
8.1.3 Implantation via Catheter or Syringes
Another minimally invasive implantation technique involves using catheters or syringes to deliver soft bioelectronics into the body. This method is particularly suitable for devices that need to be implanted deep within the body or in locations that are difficult to access through traditional surgical methods.[177]
For cardiovascular applications, devices are typically loaded into a catheter, which is navigated through the body's vascular system or other tubular structures to the target site. Once positioned, the device is deployed from the catheter, either through mechanical expansion or by being pushed out. This approach enables precise placement while minimizing tissue damage. In contrast, for applications such as injecting hydrogel-based biosensors or drug delivery systems, syringes are often employed.[178, 179] The device or material is loaded into the syringe and injected directly into the target tissue. This method is particularly effective for delivering small, injectable electronics or materials that can self-assemble or form functional structures in situ, providing a minimally invasive solution for targeted therapies or diagnostics.
For example, Liu et al. demonstrated the use of syringe injection to deliver ultrafine, flexible mesh electronics into biological and synthetic structures (Figure 11e).[180] The macroporous design allows electronics with widths over 30 times the needle diameter to be smoothly injected. Post-injection, the mesh electronics exhibit high device yield (>90%) with minimal changes in performance (≤12% impedance or conductance variation). These electronics integrate seamlessly into tissues and structures, maintaining biofunctionality while minimizing immune responses (Figure 11f).
8.1.4 In Situ Deployable Electronics
Soft implantable electronics can be designed for minimally invasive implantation through small incisions or punctures, followed by in vivo expansion or self-alignment to conform to surrounding tissue structures. These devices are typically compressed, folded, or rolled into compact configurations for insertion and then expand upon deployment using mechanisms such as shape memory polymers, hydrogels that swell in response to physiological conditions, or pre-stressed elastic materials. Additionally, self-aligning strategies leverage bioinspired designs, such as microhooks or tissue-adhesive coatings, to ensure precise positioning and secure integration with anatomical structures. These approaches minimize surgical trauma while achieving optimal alignment and functionality within the body.
Recently, Bae at al. described the use of a biodegradable and self-deployable electronic tent electrode for brain cortex interfacing.[181] This system addresses minimally invasive implantation challenges by utilizing a temporary rigid, shape-memory polymer substrate made of poly(lactide-co-ε-caprolactone)–PLGA. The material is engineered for programmable deployment via a syringe with a diameter of less than 5 mm, expanding to 200 times its initial size to cover large cortical areas (Figure 11g). The device's biocompatibility and biodegradability were validated in vivo, showing a gradual dissolution within 460 days without significant immune responses. Its electrical components, including molybdenum electrodes (≈500 nm thickness, impedance ≈15 kΩ at 1 kHz), provided stable electrocorticographic signals (signal to noise ratio, SNR ≈ 11.38 dB) for up to 14 days before degrading.
In another study, Zhang et al. developed a self-twining electrode inspired by climbing plants, utilizing a shape-memory polymer substrate (Figure 11h).[182] This electrode transitions from a flat 2D form to a 3D helical structure at body temperature (37 °C), enabling close and stable contact with target tissues. The fabrication process involved integrating gold/titanium layers onto a thin PI substrate using transfer printing and subsequent reconfiguration into a 3D shape. Electrical properties showed stable charge delivery capacity (≈9.7 mC cm−2) and low impedance (≈156 Ω at 1 kHz). Mechanically, it demonstrated remarkable flexibility, with a bending stiffness of ≈1.0 × 10⁻¹⁰ N·m2, ensuring compatibility with peripheral nerves. This design minimized mechanical mismatch and nerve stress, highlighting its potential for applications in neural interfacing and bioelectronic devices.
8.2 Chronic Fixation Process
Traditional methods for securing implants typically involve mechanical fixation techniques, such as suturing or the use of bioglue. Suturing, which involves stitching the device to surrounding tissues, is a widely employed technique for various implants, including cardiac devices and neural interfaces. Bioglue, or surgical adhesive, provides an alternative by bonding the device to tissues, offering a less invasive option particularly suited for delicate or flexible devices that could be damaged by sutures. However, these conventional methods pose significant challenges, including the risk of tissue damage, inflammation, and fibrosis at the fixation site. Suturing, in particular, can be time-intensive and unsuitable for implants requiring flexibility or movement. In this section, we explore advanced strategies for securing soft electronics to tissues with minimal damage and improved biocompatibility.
8.2.1 Passive Fixation on the Macro/Microstructures
Passive fixation involves designing implants to mechanically interlock with the natural macro- and microstructures of the target tissue.[183] This strategy relies on bioinspired features such as microhooks, barbs, or interlocking meshes that engage with the fibrous or cellular architecture of the organ (Figure 11i).[184] These designs ensure stable positioning while accommodating natural tissue movements, making them particularly suitable for dynamic environments like the heart or brain. Such design adheres to tissue without requiring adhesives or sutures, reducing the risk of inflammation or fibrosis. As an example of anchoring structure, Sunwoo et al. presented a minimally invasive approach for implanting intrinsically soft electronics into the adrenal gland using a temporary rigid shuttle that is subsequently removed after implantation.[185] The fabrication involved ultrathin PI-based probes with integrated gold electrodes encapsulated with SU-8 for biocompatibility. These probes were inserted into the adrenal gland, where an anchoring mechanism ensured stability (Figure 11j). Upon removal of the rigid shuttle, only the soft, flexible probe remained, minimizing tissue damage. Experimental results demonstrated low impedance (maintained for over 13 weeks) and high stability of recorded electrophysiological signals. The study effectively combines mechanical flexibility with chronic biocompatibility for stress monitoring in vivo.
As a representative example of self-interlocking electronics that wraps over the organ structure (Figure 11k),[186] Park et al. presented an innovative epicardial mesh using a nanocomposite of ligand-exchanged AgNWs (LE-AgNWs) and SBS, designed to mimic the mechanical and electrical properties of cardiac tissue.[187] This mesh integrates with the myocardium, enhancing biocompatibility and mechanical adaptability. The LE-AgNWs/SBS composition achieved high conductivity (11210 S cm−1) while maintaining elasticity (44.7 ± 7.5 kPa). The mesh provided synchronized electrical stimulation, reducing wall stress by 46.9 kdyne cm−2 during systole and improving systolic function with a 51% increase in fractional shortening in post-myocardial infarction rats. The device demonstrated biocompatibility with minimal inflammatory response, offering a promising approach for electromechanical cardioplasty.[188]
8.2.2 Active Fixation with Adhesive Gels
Active fixation employs adhesive gels to secure implants to tissues through physical or chemical interactions. Hydrogels with bioinspired adhesion properties, such as those mimicking mussel foot proteins, form strong yet reversible bonds with wet tissue surfaces. Alternatively, chemically cross-linkable hydrogels establish covalent bonds with tissue proteins, providing durable and stable fixation. These adhesives are particularly beneficial for soft, flexible implants as they evenly distribute stress and minimize mechanical trauma. Moreover, the tunable properties of these gels allow for the customization of adhesion strength and biodegradability to meet specific application requirements.
A notable approach to physically adhesive hydrogels (Figure 11m) was demonstrated by Chong et al., who employed a template-directed assembly technique to enhance adhesion in dynamic biological environments. Their adhesive layer, composed of activated polyacrylic acid and PEDOT:PSS, utilized chemical activation via 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide/N-hydroxysuccinimide to form covalent bonds with tissue (Figure 11n).[189] This design achieved an impressive adhesion energy of 800 J/m2 and high mechanical resilience, along with excellent biocompatibility and stable conductivity under physiological conditions, making it suitable for long-term tissue interfacing.
For chemically adhesive hydrogels (Figure 11o), Xue et al. proposed a strategy combining catechol-inspired adhesion with covalent bonding via electro-oxidized alginate-dopa (Figure 11p).[190] These hydrogels demonstrated rapid initial adhesion within seconds, followed by robust covalent linkages over hours. With adhesion strength exceeding 1268 J/m2 and stretchability greater than sevenfold, they maintained stability under cyclic strain and environmental stress. Their biocompatibility and minimal inflammatory response further confirmed their suitability for dynamic in vivo applications.
9 Practical Application of Intrinsically Soft Implantable Biosensors
Intrinsically soft implantable biosensors represent a transformative advancement in medical technology, enabling continuous monitoring and real-time feedback within the human body. Their unique combination of flexibility, stretchability, and biocompatibility allows these devices to integrate seamlessly with soft biological tissues, reducing the risk of inflammation, fibrosis, and mechanical failure. This section explores the practical applications of intrinsically soft implantable biosensors across various domains, including electrophysiology, chemical sensing, mechanical sensing, and physical sensing. It highlights specific examples of these devices in action, discusses the challenges associated with their implementation, and considers future directions for their development.
9.1 Electrophysiology Sensors
Electrophysiology sensors are designed to monitor electrical signals generated by biological tissues, such as the brain, heart, and peripheral nerves. These sensors are crucial for diagnosing and treating neurological disorders, cardiac conditions, and other health issues related to electrical activity in the body.
9.1.1 Neural Sensors
Neural interfaces rely on implantable biosensors to record and stimulate neural activity, offering applications in brain-computer interfaces, neuroprosthetics, and neurological disorder treatments (Figure 12a). Traditional neural electrodes, composed of rigid materials, induce foreign body responses due to their mechanical mismatch with soft neural tissue. Intrinsically soft neural interfaces, such as those composed of conductive hydrogels or stretchable nanocompoistes, overcome these challenges by closely mimicking the mechanical properties of neural tissue, thereby improving signal fidelity and reducing chronic inflammation.[191] For instance, Oh et al. introduced an approach utilizing 3D-printed PEDOT:PSS-ionic liquid colloidal inks for neural system applications.[59] The colloidal ink demonstrated excellent electrical conductivity (286 S cm−1) and mechanical flexibility, tolerating over 10 000 bending cycles at a radius of 1.5 mm without significant resistance changes. Its biocompatibility, achieved by removing cytotoxic ionic liquid residues via centrifugation, enabled safe in vivo use. Neural interfacing devices fabricated from colloidal ink successfully recorded optically-evoked electroencephalogram signals with high spatial resolution and signal-to-noise ratio (Figure 12b). These devices also enabled sciatic nerve stimulation at low voltages (≈60 mV), showcasing the potential for minimally invasive, long-term neural monitoring and stimulation.
In another example, Song et al. demonstrated the use of soft robotic actuation for deploying large-area, minimally invasive electrocorticography systems (Figure 12c).[192] Utilizing a flexible PDMS substrate with integrated gold microelectrodes and strain sensors, the ECoG device is inserted through a small burr hole using fluidic-driven eversion. Once deployed, the system conforms closely to cortical surfaces with minimal indentation (<2 mm). In vivo testing on a minipig recorded somatosensory evoked potentials, achieving high electrical stability and biocompatibility.
Despite these advantages, long-term stability in vivo remains a critical challenge. Soft neural interfaces are prone to material degradation, swelling, or loss of adhesion over extended implantation periods. Recent studies have focused on optimizing encapsulation materials and developing biointegrative coatings to enhance long-term performance. Furthermore, advancements in wireless power transfer and minimally invasive delivery techniques are crucial for increasing clinical feasibility.
9.1.2 Cardiac Sensors
Cardiac interfaces are essential for diagnosing and treating arrhythmias, monitoring cardiac function, and guiding interventions like ablation or pacemaker placement (Figure 12d)[193, 4] The dynamic and rhythmic contractions of the heart present unique challenges for implanted electrodes, requiring materials that can conform to the myocardium without causing mechanical irritation. Soft implantable sensors, designed with stretchable and conductive materials, seamlessly integrate with the heart's surface. These sensors provide high-fidelity electrocardiographic signals while accommodating the mechanical deformation of cardiac tissues, improving long-term performance and patient outcomes. Their biocompatible and flexible nature makes them particularly suitable for chronic cardiac monitoring and therapeutic applications. For instance, Zhou et al. introduced a bi-continuous conducting polymer hydrogel tailored for bioelectronics.[194] The material leverages PEDOT:PSS for conductivity and hydrophilic polyurethane for mechanical support, achieving a conductivity of over 11 S cm−1 and stretchability exceeding 400%. Its fracture toughness is over 3300 J m−2, maintaining stability in physiological environments. The hydrogel is printable via 3D methods, forming monolithic, soft devices for epicardial and sciatic nerve monitoring in rats (Figure 12e). Its integration ensures mild tissue response with superior long-term electrophysiological performance, including stable signal-to-noise ratios, significantly improved neural and cardiac monitoring efficacy over 28 days.
In another example done by Sunwoo et al., an approach to improve cardiac signal monitoring and treatment of ventricular tachyarrhythmias using intrinsically soft implantable electronics was presented.[195] A stretchable epicardial electrode array composed of silver nanowires and elastomers was utilized for precise multichannel mapping and targeted electrical stimulation (Figure 12f). The subthreshold stimulation protocol demonstrated an 80% success rate in arrhythmias prevention in MI-induced rabbits. This minimally invasive technique avoids myocardial excitation and electrical damage, enhancing biocompatibility and patient outcomes.
While promising, the clinical adoption of intrinsically soft cardiac interfaces requires further validation in chronic models. Stability under physiological conditions, electrode degradation, and fibrosis formation remain significant concerns. Additionally, integrating real-time data transmission and closed-loop feedback mechanisms into these devices is essential for their translation into next-generation bioelectronic therapies.
9.1.3 Muscular Sensors
Muscular interfaces enable the monitoring and stimulation of electrical activity in skeletal and smooth muscles for applications such as prosthetic control, rehabilitation, and gastrointestinal motility studies (Figure 12g).[196, 197] Traditional EMG electrodes are often limited by their bulkiness and rigidity, leading to discomfort and unreliable signal acquisition. Soft implantable electrophysiological sensors overcome these limitations by conforming to muscle tissue, allowing seamless detection and stimulation without interfering with natural movements. For example, Lee et al. employed MXene-based electrodes in fabric-based wearable electronics for real-time electrophysiological monitoring, such as ECG and EMG signals.[198] MXene electrodes, combined with cellulose nanofibers (CNF) and polycarboxylate ether (PCE), provide enhanced conductivity (24.7 Ω sq−1) and durability, maintaining stable resistance under 15% strain and over 1000 bending cycles (Figure 12h). Their optimized composition (MXene:CNF:PCE = 24:4:1) improves oxidation resistance, with R/R0 increasing modestly under humidity tests. The electrodes show comparable impedance (191 kΩ at 100 Hz) to commercial Ag/AgCl sensors, demonstrating reliable biosignal detection while addressing motion artifacts. In another example, Ji et al. demonstrated minimally invasive approach for monitoring signals in the muscular system using intrinsically soft implantable electronics.[199] The device leverages a stretchable silicon microneedle electrode array, integrated with serpentine gold interconnects encapsulated in PI, offering high biocompatibility and excellent mechanical durability under repeated strain cycles (Figure 12i). This design significantly reduces electrode-skin contact impedance, achieving a stable impedance under 0.5% variation during 45% stretching deformation. Experimental results demonstrate effective electromyogram signal acquisition with superior SNR compared to conventional wet electrodes, supporting its application in long-term muscular system monitoring and human-computer interface technologies (Figure 12j).
One of the remaining challenges in this domain is achieving stable electrode-tissue interfacing during prolonged implantation, as motion artifacts and material fatigue can degrade signal quality over time. Future research should focus on developing self-healing materials and hybrid bioelectronic architectures that integrate mechanical robustness with tissue-mimetic softness.
9.2 Physical Sensors
Physical sensors detect physical information such as pressure, strain, and temperature, providing valuable data for monitoring physiological processes. Intrinsically soft physical sensors are designed to conform to the organ surfaces, allowing for continuous and unobtrusive monitoring in a wide range of applications.[200]
9.2.1 Strain Sensors
Strain sensing is crucial for monitoring mechanical deformations in tissues, such as detecting muscle contractions, respiratory movements, or changes in vascular structures (Figure 13a). Conventional rigid sensors struggle to maintain stable contact with soft, dynamic tissues, often leading to inaccurate readings or tissue damage. Soft implantable strain sensors, composed of stretchable and biocompatible materials, conform seamlessly to the tissue, enabling precise and continuous measurement of strain without interfering with natural movement. Their flexibility and durability make them ideal for applications in rehabilitation monitoring, prosthetic control, and real-time feedback for dynamic physiological conditions. For instance, Zhang et al. presented a stretchable strain sensor system for musculoskeletal strain monitoring in live animal models (Figure 13b–d).[201] Utilizing a capacitor embedded with Au-TiO2 nanowires and silicone, the device maintains excellent durability and resolution (0.1% strain). Ex vivo, the sensor was tested on sheep tendons, demonstrating robust performance over 100 000 cycles of loading. When implanted in a sheep's medial gastrocnemius tendon, it recorded dynamic strain with a peak of 3.8% during locomotion, indicating its potential for in vivo biomechanical and clinical monitoring.
Challenges such as hysteresis, signal drift, and durability under repeated strain cycles must be addressed. Further research into fatigue-resistant nanocomposite materials could enhance the longevity and reliability of soft strain sensors.
9.2.2 Pressure Sensors
Pressure sensing is essential for monitoring intracranial pressure, blood pressure, or pressure changes in internal organs (Figure 13e). Rigid pressure sensors can induce mechanical stress on tissues, leading to discomfort or compromised functionality. Soft implantable pressure sensors, utilizing flexible membranes and piezoresistive or capacitive mechanisms, provide accurate pressure measurements while adapting to tissue movements. Pressure sensors can be applied for the tactile sensors and haptic systems. For example, Du et al. showcased an implantable, battery-free capacitive pressure sensor designed for tactile sensing (Figures 13f,g).[202] The device, using fused silica for hermetic encapsulation, features a resolution of 4.3 mN, suitable for replicating natural tactile thresholds. Implanted in a primate model, the sensor accurately monitored forces, converting capacitance changes into tactile feedback. Its wireless operation and compact size enhance its feasibility for chronic implantation in neuroprosthetic systems, demonstrating both static and dynamic sensing capabilities.
These sensors are critical for chronic monitoring in conditions like hydrocephalus, hypertension, or gastrointestinal disorders. For instance, Herbert et al. introduced a wireless vascular electronic system using printed soft sensors and multimaterial inductive stents to monitor hemodynamics in vivo.[203] Key materials included gold-coated stainless steel for the stent and elastomer-based capacitive sensors (Figure 13h). The stent showed axial stiffness comparable to commercial designs while enabling inductive wireless monitoring over 5.5 cm in air and 3.5 cm in saline. Integrated sensors exhibited high durability under cyclic pressure (up to 1000 mm Hg) and maintained functionality at a bending radius of 0.25 mm. The device successfully monitored arterial pressure, flow, and pulse in a rabbit iliac artery model, showcasing its potential in minimally invasive cardiovascular applications.
9.2.3 Temperature Sensors
Temperature sensors are used to monitor body temperature, providing valuable information for diagnosing and managing conditions such as fever, inflammation, and hypothermia (Figure 13i). Traditional temperature sensors are often rigid and uncomfortable, limiting their use in continuous monitoring. Soft temperature sensors are typically made from thermoresponsive materials, such as liquid crystals or thermistors embedded in flexible matrices. These sensors can conform to the body's surfaces and provide continuous temperature measurements. Soft temperature sensors have been used in wearable devices to monitor body temperature continuously, providing valuable data for managing conditions such as fever or hypothermia. They are also used in implantable devices to monitor temperature changes associated with inflammation or infection. For instance, Li et al. highlighted the application of intrinsically soft implantable electronics coated with a temperature-sensing hydrogel on medical catheters (Figure 13j).[204] By using a triple-network hydrogel integrated with carbon nanotube fibers as electrodes, the system achieves a high-temperature coefficient of resistance of 2.90% °C−1 and exceptional mechanical compatibility with tissues (Young's modulus ≈4.24 kPa). The coating provides early infection monitoring by detecting localized temperature changes with 0.1 °C accuracy, outperforming conventional methods in sensitivity and stability under in vivo conditions (Figure 13k). This approach significantly improves patient outcomes, such as a 90% survival rate in infection models. In other example, Kim et al. presented an implantable, soft, multimodal sensor capable of monitoring temperature and strain in vivo (Figure 13l).[205] Constructed using a tri-layered structure of Au and PI encapsulated in Ecoflex, the device demonstrated high sensitivity, stability, and biocompatibility. Applied in neural systems, the sensor exhibited minimal interference, allowing accurate decoupling of temperature and strain signals during nerve monitoring. Continuous monitoring in animal models revealed precise detection of strain increases (up to 3%) and temperature changes (0.6 °C), emphasizing its potential for clinical scenarios requiring long-term, minimally invasive monitoring (Figure 13m). Recently, Madhvapathy et al. The study presents miniaturized wireless temperature sensors for monitoring intestinal inflammation in a Crohn's disease-like mouse model (Figure 13n).[206] The sensors demonstrated precise monitoring with a resolution of 0.01 °C over 4 months (Figure 13o). Implantation against the abdominal muscle with smooth silicone surfaces minimized immune response. Results linked intestinal temperature fluctuations to inflammatory cytokine levels, providing real-time insights into disease progression and offering a potential tool for early intervention in inflammatory bowel diseases.
9.3 Chemical Sensors
Chemical sensors detect specific biochemical markers or analytes in the body, providing valuable information for diagnosing and monitoring various medical conditions. Intrinsically soft chemical sensors are particularly advantageous for continuous, real-time monitoring in dynamic environments such as blood vessels or interstitial fluid.
9.3.1 pH Sensing
Monitoring pH levels in tissues and organs is critical for detecting pathological conditions such as ischemia, acidosis, or infections (Figure 14a). Traditional rigid pH sensors suffer from poor biocompatibility and long-term instability due to biofouling. Soft implantable pH sensors, designed with flexible hydrogel matrices or ion-selective membranes, conform seamlessly to the body's complex structures. These sensors enable continuous monitoring of pH changes in real-time, providing valuable insights for diagnosing diseases and assessing the local biochemical environment, such as acidification in tumors or ischemic regions. Their biocompatibility and adaptability to dynamic tissues ensure minimal disruption to the surrounding environment while maintaining long-term functionality. For example, Li et al. introduced a bioresorbable, pH-responsive hydrogel sensor integrated with a wireless inductor-capacitor circuit.[207] The hydrogel, composed of PDPAEMA-PEGDA, exhibits volumetric expansion in acidic environments (pH < 6.3), which alters the inductance and resonates with a readable frequency shift. Animal trials demonstrated precise pH monitoring during gastric leaks with minimal biocompatibility issues over a 7-day functional period (Figure 14b,c). In another example, Nam et al. presented a minimally invasive approach using an intrinsically soft implantable microfiber, comprising a liquid metal core and a multifunctional nanocomposite shell with IrO2 nanoparticles (Figure 14d,e).[133] The microfiber detects pH levels and records electrophysiological signals in vivo. Its strain-insensitive conductivity (3.4 × 10⁴ S cm−1) and low impedance enable seamless tissue integration and reliable signal monitoring. In gastric applications, it provides precise pH while maintaining biocompatibility, demonstrating potential for advanced bioelectronic interfaces (Figure 14f,g).
While soft pH sensors have demonstrated improved in vivo performance, further research is needed to enhance their long-term calibration stability and drift resistance. Future developments should focus on self-calibrating systems that compensate for environmental fluctuations.
9.3.2 Ion Sensing
Ion sensors play a vital role in monitoring critical electrolyte levels, such as potassium, sodium, calcium, and chloride, which are essential for maintaining homeostasis and assessing organ function (Figure 14h). Abnormal ion concentrations are associated with various conditions, including heart arrhythmias, kidney dysfunction, and neurological disorders. Traditional rigid ion-selective electrodes face challenges in achieving stable contact with soft tissues. Soft implantable ion sensors, incorporating stretchable ion-selective membranes and conductive polymers, address these limitations by providing accurate and continuous ion monitoring in physiological environments. These sensors are particularly valuable for cardiac applications, where real-time potassium and calcium monitoring can help prevent life-threatening arrhythmias, and for renal function assessment in patients with electrolyte imbalances. For instance, Yang et al. presented a flexible multifunctional electrode using a CNT array to monitor both electrocorticography signals and ion concentrations (Ca2⁺, K⁺, Na⁺) in the cerebral cortex (Figure 14i).[208] The CNT array provides high electrical conductivity, low impedance, and stability under physiological conditions. Its excellent mechanical and ionic response properties enable real-time monitoring in animal models, demonstrating long-term biocompatibility and dynamic signal stability (Figure 14j,k).
However, long-term biofouling and signal drift present major challenges in chronic ion sensing applications. Recent efforts have explored antifouling coatings, zwitterionic hydrogels, and biomimetic ion channels to improve stability. Addressing these challenges will be crucial for translating soft ion sensors into real-world medical applications.
9.3.3 Biomolecular Sensors
Implantable biomolecular sensors are essential for detecting biomolecules such as glucose, proteins, and enzymes, which serve as biomarkers for various diseases (Figure 14l). Glucose sensors, for instance, are critical for managing diabetes, while protein and enzyme sensors can aid in diagnosing cancer or monitoring metabolic activity.[209, 210] Conventional sensors often struggle with biocompatibility and long-term stability. Soft implantable biomolecular sensors, utilizing enzyme-functionalized hydrogels or nanocomposite coatings, provide a flexible and minimally invasive solution. These sensors offer high sensitivity and specificity, enabling real-time detection of target biomolecules directly at the site of interest. For example, glucose sensors integrated with soft, biocompatible matrices can continuously monitor blood glucose levels, reducing the need for frequent invasive testing. These advances ensure reliable performance and improved patient outcomes in chronic disease management and personalized medicine. For the specific example, Li et al. demonstrated the application of intrinsically soft implantable electronics as biomolecular sensors for in vivo monitoring.[211] The study uses highly flexible, biocompatible materials like PDMS-encapsulated CNTs for high conductivity and mechanical robustness (Figure 14m). The sensor achieved precise real-time detection of biomarkers with over 98% accuracy, showing exceptional stability under physiological conditions. Its integration into live tissue enabled continuous monitoring, offering promising advancements in diagnostic and therapeutic applications within living systems (Figure 14n).
Despite these advances, long-term enzyme stability and sensor degradation remain limiting factors. Strategies such as enzyme encapsulation within nanoporous matrices and synthetic biomimetic receptors are being explored to extend operational lifespan. Additionally, integrating these sensors with soft wireless power and data transmission modules could enable real-time, remote monitoring of metabolic biomarkers.
10 Current Limitations and Future Prospects
The field of intrinsically soft implantable biosensors has advanced significantly, offering transformative solutions for continuous monitoring and therapeutic interventions. Despite these strides, several challenges must be overcome to fully realize their potential. This chapter outlines the current limitations in material design, biocompatibility, power, and data transmission and explores future prospects for these technologies.
10.1 Current Limitations
10.1.1 Material and Mechanical Challenges
The primary hurdle in developing soft biosensors lies in balancing mechanical compliance and electrical functionality. Soft materials like hydrogels and elastomers often degrade under continuous mechanical stress, such as cardiac or muscular movement, leading to reduced durability. Conductive polymers and nanomaterials, while enhancing electrical performance, face challenges with integration into soft matrices, where conductivity can deteriorate in the harsh physiological environment. Additionally, combining multiple components—sensors, power sources, and communication modules—into compact, flexible devices remains complex, as does ensuring robust encapsulation against corrosive bodily fluids.
10.1.2 Biocompatibility and Stability
While soft materials are inherently more biocompatible than rigid counterparts, long-term implantation introduces challenges such as chronic inflammation, fibrosis, and biofouling. The accumulation of biological materials can obstruct sensing elements, impairing functionality. Degradable materials, often used for temporary implants, require precise control of degradation rates to ensure sustained functionality without premature failure.
10.1.3 Clinical Feasibility and Regulatory Considerations
While significant progress has been made in functionalizing soft implantable biosensors for improved biocompatibility, clinical translation requires addressing practical feasibility and regulatory hurdles. Regulatory agencies such as the FDA and EMA impose stringent biocompatibility requirements, including cytotoxicity, hemocompatibility, and long-term stability assessments.
One major challenge in clinical implementation is the long-term mechanical and chemical stability of these materials under physiological conditions. For instance, encapsulation layers that degrade over time may lead to exposure of underlying materials, triggering immune responses. Additionally, fibrosis and chronic inflammation around implants remain key barriers to long-term functionality.
To bridge the gap between preclinical research and clinical applications, systematic in vivo studies, including large-animal models and human trials, are necessary. Future research should focus on developing standardized testing protocols for evaluating the long-term stability, biointegration, and safety of these materials in clinical settings. Furthermore, interdisciplinary collaboration between materials scientists, biomedical engineers, and clinicians will be essential in refining these strategies for real-world applications.
10.1.4 Power and Data Transmission
Powering and communicating with implantable devices remain a critical challenge. Batteries, while effective, are bulky, limited in lifespan, and require replacement surgeries. Emerging energy harvesting techniques, such as converting body heat or motion into electricity, are promising but face integration hurdles.[212] Similarly, reliable wireless data transmission through biological tissues demands advancements in signal efficiency and security to safeguard patient data.
10.2 Future Prospects
Future research in soft implantable biosensors should focus on developing advanced materials that integrate mechanical softness, durability, and multifunctionality. Hybrid materials combining conductive hydrogels with metallic nanostructures or bioactive polymers could enhance both electrical stability and biointegration, ensuring long-term performance in physiological environments. Additionally, the incorporation of bioresorbable materials that gradually degrade after fulfilling their sensing function may lead to transient implantable biosensors, eliminating the need for device retrieval and reducing long-term complications.
Beyond material engineering, the advancement of smart and adaptive biosensors presents a promising direction. Self-healing properties, where damaged materials autonomously repair themselves, could significantly extend device lifespan, mitigating mechanical failures that arise from chronic implantation. Moreover, biosensors capable of dynamically responding to physiological changes, such as swelling or temperature fluctuations, would improve their long-term reliability. Integrating machine learning algorithms into biosensing systems could further enhance diagnostic accuracy by enabling real-time adaptation of sensing parameters, improving both signal processing and automated feedback control.
Despite these innovations, the clinical translation of soft implantable biosensors remains a major challenge. While preclinical studies have demonstrated promising results, many devices have yet to undergo large-scale animal studies or human trials necessary to assess their long-term functionality, safety, and stability under real-world conditions. Additionally, regulatory approval pathways for novel soft bioelectronics remain complex, requiring standardized protocols for material characterization, reliability testing, and biocompatibility evaluation. Without clear regulatory guidelines, the transition from laboratory research to clinical application remains slow and uncertain.
Another crucial aspect of future development lies in the integration of wireless, closed-loop sensing and therapeutic systems. The next generation of soft bioelectronics should not only record physiological signals but also provide real-time feedback to modulate biological activity, such as neuromodulation for neurological disorders or cardiac pacing for arrhythmias. Coupling these systems with wearable or cloud-based monitoring platforms will enable continuous, remote healthcare management, allowing for more personalized and proactive patient care. However, ensuring stable, long-range wireless communication while maintaining the mechanical flexibility and biocompatibility of these devices remains a technical hurdle that must be addressed.
As soft implantable biosensors approach clinical use, ethical and societal considerations must also be carefully examined. Long-term implantation raises concerns about data security, patient autonomy, and equitable access to emerging technologies. Establishing transparent frameworks for patient consent, privacy protection, and responsible technology deployment will be crucial in ensuring the ethical integration of these devices into medical practice. Furthermore, considerations surrounding cost and accessibility must be addressed to prevent disparities in healthcare access, ensuring that these innovations benefit a broad patient population rather than a select few.
By addressing these challenges and advancing research in material engineering, adaptive electronics, wireless integration, clinical validation, and ethical considerations, the field of soft implantable biosensors can move closer to achieving widespread clinical adoption. Continued interdisciplinary collaboration among material scientists, engineers, medical professionals, and regulatory bodies will be essential in translating these innovations into practical, life-enhancing solutions.
Acknowledgements
S.H.L. and H.J.L. contributed equally to this work. This research was supported by the Institute for Basic Science (IBS-R006-A1). This research was also supported by the National Research Foundation of Korea (NRF) grant funded by the Korean government (2021R1C1C2004400).
Conflict of Interest
The authors declare no conflict of interest.
Biographies
D.-H. Kim obtained his B.S. and M.S. degrees in chemical engineering from Seoul National University, Korea, in 2000 and 2002, respectively. He received his Ph.D. degree in materials science and engineering from the University of Illinois at Urbana–Champaign in 2009. From 2009 to 2011, he was a postdoctoral research associate at the University of Illinois. He joined Seoul National University in 2011 and is currently a professor in the School of Chemical and Biological Engineering of Seoul National University. He has been serving as an associate director of the Center for Nanoparticle Research of the Institute for Basic Science (IBS) since 2017. He has been focusing on the research of nanomaterials and deformable devices and their application to bio-integrated and bioinspired electronics. He has been recognized with several awards including the George Smith Award (2009), TR 35 Award (2011), Hong Jin-ki Creative Award (2015), SCEJ Award (2016), and Korea Young Scientist Award (2017). He was also selected as one of the highly cited researchers by Clarivate Analytics in 2018–2024.
H. J. Kim received her Ph.D. degree in electrical and computer engineering from Korea University, Seoul, Korea, in 2018. From 2018 to 2021, she worked as a research professor in the Department of Clinical Pharmacology and Therapeutics at the College of Medicine, Kyung Hee University, Seoul, Republic of Korea. Between 2021 and 2024, she held positions as a junior and senior researcher at the Institute of Chemical Process (ICP) at Seoul National University and at the Center for Nanoparticle Research, Institute for Basic Science (IBS), both in Seoul, Republic of Korea. In 2024, she joined the Department of Biomedical Engineering at Yonsei University, Wonju, Republic of Korea, as an assistant professor. Her current research focuses on biosensors and soft bioelectronics aimed at translational research.
S-H Sunwoo received his B.S. (2015) and M.S. (2018) degrees from the Department of Biomedical Engineering at Sungkyunkwan University. He obtained his Ph.D. (2022) degree from the Department of Chemical and Biological Engineering at Seoul National University. Since he joined the faculty of the Department of Chemical Engineering at Kumoh National Institute of Technology in 2024, he has focused on soft bioelectronics for diagnostic and therapeutical purposes.
