Scalable and stretchable 1D multifunctional fibers for multimodal sensing and stimulation

Domain: biology
Published: 2026-03-11
Editorial outcome: major_revision

Abstract
One-dimensional (1D) multifunctional fibers have garnered significant attention due to their advantageous geometry properties, which allows conformal interfacing with soft biological tissues and efficient charge transport. Here, we developed a solution-deposition strategy for the scalable and cost-effective fabrication of stretchable liquid metal fibers integrated with electrochemically stable, tissue-interfacing electrodes, thereby enabling the realization of stretchable multifunctional fibers. This fiber seamlessly combines electrodes and conductive pathways into a single structure, enabling versatile applications such as electrophysiological signal sensing, in vivo nerve stimulation, and wireless energy transmission. The multifunctional fiber demonstrates significantly improved electrical performance under strain, maintaining conductivity during stretching and bending, and exhibits lower impedance and higher signal stability, particularly during physiological monitoring and electrical stimulation. The fiber’s excellent biocompatibility and mechanical compliance makes it well suited for wearable systems and long-term biomedical applications, offering a robust platform for next generation 1D bioelectronics.

Introduction
Stretchable one-dimensional (1D) multifunctional fibers exhibit superior flexibility, isotropic deformability, breathability, and lightweight characteristics, which are promising for emerging areas such as wearable systems, smart sutures, and minimally invasive medical devices1–6. Although textile electronics have achieved considerable success in wearable applications, single fibers offer superior characteristics for particular applications7–16. Their inherent flexibility, stable conductivity, sensing capabilities, and reliable signal transmission in dynamic environments make them exceptionally well-suited to integrate biological tissues as functional platforms17–19. Electrical conductivity is a fundamental requirement for multifunctional fibers, and a widely adopted strategy for developing stretchable 1D conductors involves hybridizing or coating methodologies that synergistically integrate elastic polymeric substrates with conductive nanomaterials20–24. The electrical contact between conductive fillers is pivotal in determining conductivity, regardless of the material selected. However, during stretching, the mechanical property mismatch between the conductive material and the matrix, coupled with the separation of fillers, leads to a marked deterioration in conductivity25,26. Moreover, stretchable 1D conductors must exhibit dynamic conductivity and ensure stable electrochemical performance to facilitate prolonged and reliable dynamic monitoring within the in vivo environment27. It remains challenging to fabricate stretchable fibers that meet the demands of conductivity, scalability, and stable electrochemical performance.

In recent years, liquid metals have been recognized as a promising partial solution to this challenge. They have distinctive properties, including a low gauge factor, exceptional deformability with zero modulus, and high electrical conductivity28–31. However, liquid metals exhibit inherent limitations that must be addressed to fully harness their potential for developing stretchable 1D multifunctional fibers. Firstly, liquid metals possess high surface tension, which impedes their direct adhesion to elastic polymers32,33. To address this issue, a commonly employed method involves incorporating a significant amount of gallium oxide into the liquid metal to reduce surface tension34,35. Nevertheless, such approach provides limited control the microscopic interfacial contact between the polymer and the liquid metal, making it difficult to achieve a uniform distribution. Another approach utilizes the direct fabrication of liquid metal fibers through injection and thermal drawing techniques, which are intrinsically constrained by the size of the fiber dimensions, typically yielding a minimum diameter of 250 µm36–38. Furthermore, these liquid metal fibers cannot directly interface with biological tissues or fluids; instead, they can only indirectly sense biomechanical or thermal signals through deformation or temperature changes transmitted via surrounding elastomer layers. The surrounding biofluid contains water and ions that rapidly react with gallium, thereby inhibiting stable and reliable operation for long-term in vivo applications29,39. Although various solution-based methods for preparing liquid metals have been reported, the large-scale fabrication of a single liquid metal fiber remains unexplored40–43. The main challenge arises from the structural differences between films and fibers: large-area films possess abundant reactive interfaces that readily interact with liquid metals, whereas single fibers have a much smaller surface area and thus require higher-quality electrochemical coatings to achieve uniform and stable encapsulation. This constraint significantly limits their performance in implantable electrophysiological signal detection and electrical stimulation of biological tissues. Collectively, These practical challenges pose major barriers to liquid metal-based 1D bioelectronics.

In this study, we developed a scalable fabrication strategy for stretchable, multifunctional 1D bioelectronics fibers. We first established single liquid-metal fibers to enable efficient signal transmission and subsequently constructed PEDOT:PSS inert electrodes on selected regions of the fibers, thereby enabling both sensing and electrical stimulation capabilities (Fig. 1). Furthermore, by twisting and assembling multiple encapsulated liquid-metal electrodes, a multichannel sensing system can be realized (Fig. 1b). Specifically, a polydopamine (PDA) layer was formed via the self-polymerization of dopamine solution, providing a universal surface functionalization that facilitates electroless deposition of a copper (Cu) layer onto the fibers (Fig. 1c). The Cu layer subsequently serves as a stabilizing interface through reactive wetting with the liquid metal (Fig. 1d). Since this method relies on liquid-phase chemical reactions to stabilize the adhesion of liquid metal on the fiber surface, it overcomes the size constraints of conventional fabrication approaches (e.g., infusion and thermal extrusion), achieving a fiber diameter as small as 50–100 µm. By segmentally encapsulating the liquid-metal fibers with materials of distinct functionalities, such as insulating SEBS elastomers or stretchable, chemically stable conductive carbon nanocomposites, this process yields mechanically robust and electrochemically stable microelectrodes suitable for long-term applications (Fig. 1e–g). The engineered stretchable 1D multifunctional fibers exhibit excellent electrical conductivity and maintain stable performance under mechanical deformation, as demonstrated by their ability to continuously power commercial light emitting diodes (LEDs) during stretching (Fig. 1h, i). We further introduced an inert CNT/PEDOT:PSS electrode layer, allowing the exposed regions to directly interface with biofluids and perform stable sensing functionality. This bilayer electrode architecture, composed of liquid metal and inert functional layer, integrates high conductivity with reliable sensing performance under complex biological conditions (Fig. 1j). Applied to multimodal electrophysiological monitoring, these stretchable bioelectronic fibers effectively suppress motion artifacts during dynamic activities. When combined with machine learning algorithms, the platform achieved an action-recognition of 99.5%. Additionally, leveraging the fibers’ high electrochemical stability and biocompatibility, we performed precise electrical stimulation of the sciatic nerve in animal models, successfully modulating muscle movements. Overall, these results highlight the platform's capability for bidirectional bioelectronic interface with broad potential for diverse biomedical and wearable applications.

Results
The fabrication of the stretchable multifunctional fiber begins with the formation of a stretchable conductive layer. To produce continuous one-dimensional liquid metal layer, we employed a chemical solution deposition process based on the alloying reaction between gallium-based liquid metal and copper, as shown in Fig. 1c. In contrast to conventional methods (Fig. 1a), which typically require the repetitive application of oxidized liquid metal to ensure adhesion to the fiber, this solution-based manufacturing process enables continuous, scalable production while ensuring a more uniform distribution of liquid metal across the fiber surface. The fabrication begins with an ultrathin polyurethane (PU) fiber core, which provides intrinsic mechanical stretchability to the electronic system (Fig. S9). A PDA adhesive layer is applied to the fiber surface, creating abundant and uniformly distributed active sites for silver nanoparticle (AgNP) deposition. Successive copper layers are subsequently deposited onto the AgNP-coated surface, and all these steps are performed using low-cost, scalable solution-based methods. In the final step, liquid metal is introduced in dilute hydrochloric acid, selectively alloying with the copper to form a continuous, high-conductivity pathway along the fiber. As schematically illustrated in Fig. 2a, the stretchable, high-integration, multifunctional conductive fiber fabrication relies on a rapid alloying reaction between the Cu coating and the liquid metal. This alloying process ensures robust adhesion and efficient electronic transport pathways across the fiber structure (Fig. 2b). To evaluate the wetting behavior critical for the alloying process, the contact angle of a liquid metal droplet was compared on two substrates: the untreated PU substrate and the chemically deposited Cu substrate (Fig. 2c). The contact angle on the PU surface is relatively high at 119.7°, indicating poor wettability, whereas the Cu-coated surface significantly reduces the contact angle to 36.5°, facilitating improved wetting and interaction with liquid metal. This enhanced wettability is critical for ensuring the uniform alloying process during fabrication.

The uniformity and compositional integrity of the electronic fiber at various stages of the layer-by-layer fabrication process were further validated using SEM imaging and elemental mapping (Fig. 2d). The elemental distribution images clearly illustrate the growth of AgNPs on the PDA-coated fiber, the subsequent Cu surface modification on the Ag fiber, and the uniform spatial distribution of In after the final alloying reaction between Cu and liquid metal. The multilayer fabrication process resulted in only a minimal increase in the overall fiber diameter, essentially adding a thin surface layer of approximately 2–3 µm (Fig. S2). Moreover, the corresponding stress-strain behavior was not significantly affected (Fig. S6). In contrast, the electrical conductivity changed markedly at each fabrication stage, as shown in Fig. S7. The initial Ag layer failed to form a stable conductive network, resulting in negligible conductivity. The subsequent Cu coating achieved a low conductivity of 0.1133 S cm⁻¹. Finally, after reaction with the liquid metal, the conductivity dramatically increased to 39,623 S cm⁻¹. The thickness and deposition amount of the internal layers (PDA, AgNW, and Cu) were found to be highly dependent on the deposition time. As summarized in Fig. S8, the optimized parameters are as follows: a 14-h PDA deposition formed a uniform and continuous adhesive layer, providing an ideal foundation for subsequent metallization. A 3-h AgNP deposition achieved sufficient surface coverage to act as an effective conductive activation layer, while a 30-min copper electroplating produced a continuous and dense metallic film that served as an excellent conductive substrate for the subsequent Galinstan coating. Collectively, these optimized parameters ensure ideal, full-surface coverage of the final liquid metal layer. To uncover the chemical composition and surface morphology, X-ray photoelectron spectroscopy (XPS) analysis was performed to elucidate the underlying mechanisms of the solution-processed liquid metal patterning method (Fig. 2e). After treatment in silver nitrate solution, the XPS spectrum (TPU/PDA/Ag) exhibits peaks at 374 eV (Ag3d 3/2) and 368 eV (Ag3d 5/2), verifying the formation of Ag seed nanoparticles27. These nanoparticles, which are sparsely distributed on the PDA film, have sizes in the tens of nanometers. The AgNP seeds catalyze the subsequent chemical deposition of the Cu film, as evidenced by the appearance of Cu2p1/2 (952.6 eV) and Cu2p3/2 (932.8 eV) peaks in the XPS spectrum (TPU/PDA/Ag/Cu)44. The structural evolution of the fiber was further analyzed through X-ray diffraction (XRD), as shown in Fig. 2f. The XRD spectrum of the chemically deposited Cu layer exhibits two broad peaks at 43.3° and 50.5°, corresponding to fine-grained crystalline Cu. Following the deposition of the liquid metal, the XRD spectrum reveals a series of sharp crystalline peaks, attributed to the formation of CuGa and CuGa2 intermetallic compounds45,46. The alloying reaction occurring at the Cu-liquid metal interface is a key driving force that significantly enhances the wettability of the non-oxidized liquid metal47. This improved wettability is crucial for achieving a uniform and continuous conductive pathway. The Energy Dispersive X-ray Spectroscopy (EDS) spectrum acquired from the selected area clearly shows the characteristic peaks of Ag, Ga, In, and Sn. Ga is identified as the dominant phase, with the significant presence of Cu being consistent with the formation of a Cu-Ga intermetallic alloy (Fig. S4).

The uniformity and quality of the liquid metal coating are crucial for ensuring the reliable performance of fiber electrodes. This deposition process is influenced by several key processing parameters. To establish a robust and reproducible process window, we systematically investigated the effects of drawing speed, fiber modulus, ambient temperature, and the volume of Galinstan droplets on the final coating morphology. In Fig. S3, at speeds approaching 20 cm/s, the reaction between the liquid metal and copper becomes kinetically limited, resulting in non-uniform coverage and a defective interface. This finding allows us to define a precise operating window for high-speed coating. In contrast, the Young’s modulus of the elastic fiber substrate, ambient temperature, and the volume of liquid metal exhibited unexpectedly minimal influence within the tested ranges, indicating broad material compatibility of our approach. Furthermore, cross-sectional SEM imaging together with elemental analysis demonstrates the robust adhesion of liquid metal to the TPU fibers (Fig. S5). A complementary video (Supplementary Video 1) was recorded to validate the superior durability and adhesion of our conductive fiber under dynamic mechanical stress. The footage shows the sensor undergoing repeated stretching and bending cycles. Notably, the metallic coating retained its structural integrity throughout, with no evidence of cracking, delamination, or interfacial detachment. The high throughput of the continuous melt-spinning and all-solution processes enabled the successful production of 50-meter-long TPU fibers and corresponding liquid metal-coated fibers, demonstrating their potential for scalable manufacturing (Fig. S10).

The encapsulation layer is critical in the multifunctional fiber’s water resistance and stable electrical performance. As shown in Fig. 2g, the elastic polymer encapsulation can protect the liquid metal fiber while maintaining efficient electron transport through the conductive pathway. We use the feasible drop-casting process to encapsulate the liquid metal fiber. Figure 2h illustrates the encapsulation process, where polymer droplets are uniformly coated onto the fiber surface to seal the liquid metal core. This encapsulation enhances the durability of the fiber by protecting it from oxidation, corrosion, or current leakage when contacting bio-tissues. Stress-strain measurements were conducted to evaluate the mechanical performance of the encapsulated liquid metal fiber, as shown in Fig. 2i. The fiber demonstrates an impressive stretchability of up to 250%, withstanding significant tensile strain without mechanical failure. Electrical performance under mechanical deformation was further assessed by monitoring the normalized resistance as a function of uniaxial tensile strain (Fig. 2j). The resistance increases by two at 100% strain and by a factor of 9 at 200% strain. These results indicate that the liquid metal fiber maintains low resistance and stable conductivity under high tensile strain, demonstrating its potential for dynamic applications requiring flexibility and durability. The chemical stability of Gallium-based liquid metal is known to be unsatisfactory, especially under physiological conditions. As shown in the inset of Fig. 2k, significant chemical corrosion of the liquid metal appeared after a few hours of immersion in phosphate-buffered saline (PBS) solution. Correspondingly, the resistance increases dramatically by over 100 times within 4.5 h. In contrast, when coated with a 25 μm thick SEBS (styrene-ethylene-butylene-styrene) elastomer layer, the resistance of the liquid metal fiber remains nearly unchanged after prolonged storage in PBS solution, further confirming its ability to maintain stable electrical performance. The thin elastomer encapsulation layer is therefore indispensable for ensuring the long-term stability of liquid metal features in practical applications. By controlling the movement of the polymer solution droplet along the liquid metal fiber, the liquid metal fiber can be selectively encapsulated with a specific section exposed (Supplementary Video 2). As shown in Fig. 2l, we conducted the continuous resistance measurements along the fiber to validate the selectively encapsulated liquid metal fiber. The encapsulated fiber exhibits a consistent voltage distribution throughout its length, indicating uniform conductivity and reliable performance. This programmable encapsulation further enhances the multifunctionality of the conductive fiber, making it suitable for a wider range of applications requiring precise electronic control.

Building upon the successful fabrication of the liquid metal layer, the stretchable multifunctional fiber is further functionalized with an electrochemically inert electrode layer that is fully compatible with physiological environments (Fig. 3a). Stretchable nanocomposite materials represent an attractive platform for designing electrode with tailored properties. Multi-walled carbon nanotubes (MWCNTs) were selected as conductive fillers due to their high electrical conductivity and excellent electrochemical stability. By dispersing MWCNTs into a SEBS (styrene-ethylene-butylene-styrene) solution to form a viscous ink, patterned conductive nanocomposites could be feasibly deposited via a mask-spray coating process. To further reduce impedance and enable high-fidelity interfacing between the electrode and biological tissues, a layer of poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) (PEDOT:PSS) was electrochemically deposited onto the MWCNT electrodes (Fig. 3b). Surface morphology and elemental composition of the PEDOT:PSS-coated fibers were analyzed using SEM and elemental mapping (Fig. 3c). The SEM images revealed a uniform PEDOT:PSS deposition on the nanocomposite electrode, and the inset confirms the presence and homogeneous distribution of sulfur, indicating the formation of the PEDOT:PSS layer. Volumetric capacitance and cathodic charge storage capacity, as key indicators of the PEDOT:PSS coating, were evaluated for electrodes subjected to varying numbers of deposition cycles (Figs. S11 and S12). Their electrochemical behavior was assessed in PBS solution, revealing a clear saturation effect: the electrochemical capacity plateaued once the number of PEDOT:PSS deposition cycles reached 30. Considering this saturation behavior alongside a comprehensive performance evaluation, 30 electrochemical deposition cycles were selected as the optimal condition. This uniform coating enhances high-fidelity signal recording while maintaining mechanical flexibility. The electrochemical stability of the conductive nanocomposites was further evaluated through linear sweep voltammetry (Fig. 3d). The results demonstrate a broad electrochemical stability window spanning from −1.85 to 2.15 V, confirming the reliability of the CNT/PEDOT: PSS electrode under physiological conditions. Conversely, when subjected to applied positive voltage, the liquid metal fiber exhibits performance degradation. This degradation process can also release high concentrations of gallium and indium ions, which pose significant cytotoxicity risks. It highlights the necessity of employing an electrochemical deposition process to enhance its stability.

The cyclic voltammetry (CV) performance of the fiber electrodes is shown in Fig. 3e. The pristine nanocomposite electrodes exhibit minimal current response over the entire voltage range during the 1st and 40th cycles, indicating poor electrochemical activity and limited charge storage capacity. In contrast, the PEDOT:PSS-coated electrodes show significantly larger and more stable current responses across multiple cycles, highlighting the enhanced charge transport properties and electrochemical activity introduced by the PEDOT: PSS layer. Furthermore, the consistent CV curves over repeated cycles demonstrate excellent cycling stability, making the modified electrodes ideal for applications requiring reliable and efficient charge transfer, such as bioelectronic sensing and stimulation. The mechanical performance of the fiber electrodes under uniaxial tensile strain was evaluated by monitoring the normalized resistance as a function of strain (Fig. 3f). The normalized resistance increases to 7.2 at 50% strain and 16.9 at 100% strain. This excellent stretchability is attributed to the electrodes forming a structurally continuous and intrinsically deformable conductive network within the SEBS elastomer. Such mechanical stability ensures reliable electrical performance in dynamic applications requiring repeated deformations.

Electrochemical impedance spectroscopy (EIS) was used to evaluate the electrical performance of the fabricated fiber electrodes. As shown in Fig. 3g, the bare nanocomposite electrode impedance magnitude is relatively high, which can be attributed to the limited active surface area of the exposed MWCNTs. This high impedance poses a significant challenge for high-fidelity recordings due to signal attenuation and distortion. Electrochemically depositing a PEDOT:PSS layer onto the fiber electrodes reduced the impedance by two orders of magnitude. The phase angle analysis further reveals a phase angle approaching 0°, indicating a dominant resistive component suitable for electrophysiological recording and stimulation. The impedance stability of the fiber electrodes was assessed under mechanical strain (Fig. 3h). The PEDOT:PSS layer exhibited relatively stable interfacial impedance during stretching, which can be attributed to the volumetric charge storage mechanism36. The fiber electrodes underwent 1000 stretching cycles at 30% strain to evaluate durability. The comparison of impedance at 1 kHz before and after cycling (Fig. 3i) shows negligible changes, confirming the outstanding mechanical and electrical stability of the PEDOT:PSS-coated fibers. To ensure the reliability of the fiber electrodes in physiological environments, we evaluated the long-term stability of the PEDOT:PSS functional layer in simulated body fluid (phosphate-buffered saline, PBS, pH 7.4). Throughout the immersion period, no significant increase or fluctuation in impedance was observed, indicating that the interfacial properties between the coating and the electrolyte remained intact (Fig. S13).

The fiber-like structural characteristics and scalable manufacturing capabilities of stretchable multifunctional 1D fibers offer high potential for integration with textiles through embroidery techniques. In this study, we embroidered these fibers onto fabrics to fabricate radio-frequency (RF) antennas for wireless power transmission. As shown in Fig. 4a and Supplementary Video 3, conductive liquid metal fibers are stitched into flexible electronic textiles using a programmable embroidery process, creating textile-based inductors with complex, stable conductive pathways. The inset highlights the uniform structure of the encapsulated conductive fibers, which ensures stable and reliable performance under mechanical deformation. Figure 4b demonstrates the flexibility and adaptability of the electronic textile under various mechanical deformations, including bending, twisting, and stretching. Despite these dynamic deformations, the textile maintains its structural integrity. The electrical stability of the textile inductors under mechanical stress was evaluated by monitoring resistance changes during bending, twisting, and hammer impacts (Fig. 4c). The results show that the resistance remains constant under these conditions, verifying the excellent mechanical stability of the embroidered conductive fibers. This stability ensures consistent performance in real-world applications, even under repeated dynamic mechanical stress.

In practical applications, electronic textiles integrated with conductive fibers conform well to curved biological surfaces. As shown in Fig. 4d, the textile closely conforms to the palm, demonstrating its potential for integration into wearable bioelectronic systems. Moreover, due to the excellent waterproof properties of the encapsulated conductive fibers, the textile successfully transmits energy. It maintains stable illumination of multiple light-emitting diodes (LEDs) even in the presence of water droplets (Fig. 4e). This highlights the waterproof characteristics of the textile inductors and their reliable wireless energy transmission performance under complex environmental conditions. To compare the wireless energy transmission efficiency of the liquid metal fiber inductors with conventional copper wire inductors, the quality factor (Q) was measured across a range of frequencies (Fig. 4f). The results show that the Q-factor of the liquid metal fiber inductors is comparable to that of copper wire inductors and performs exceptionally well at higher frequencies, demonstrating their capability for efficient energy transmission. The mechanical flexibility and durability of the electronic textiles were further assessed through cyclic bending tests. Figure 4g presents the resistance as a function of bending cycles, where the copper wire fractures after around 170 cycles. In contrast, the liquid metal conductive fibers exhibit superior yield strain, maintaining minimal resistance variation even after 400 bending cycles, significantly outperforming commercial copper wires’ durability.

Additionally, Fig. 4h compares the resistance during the first and final bending cycles, showing negligible changes, further confirming the mechanical and electrical stability of the embroidered conductive fibers over extended use. The textile resistance remains constant during each bending cycle, underscoring its consistent performance. These characteristics make the embroidered textile conductive traces well-suited for next-generation wearable electronic systems, enabling stable operation under mechanical deformation and complex environmental conditions while facilitating efficient wireless energy transmission. These features make the embroidered textile inductors highly suitable for next-generation wearable electronic systems, capable of stable operation under mechanical deformation and complex environmental conditions while enabling efficient wireless energy transmission.

The multifunctional fiber was further utilized for electrophysiological monitoring, enabling high-fidelity and stable recordings of electrocardiogram (ECG) and electromyogram (EMG) signals. It exhibited excellent performance and robustness under various physiological and environmental conditions. As shown in Fig. 5a, the multifunctional fibers were wrapped around volunteer’s wrists, with the electrodes maintaining tight contact with the skin for ECG signal acquisition. By integrating the fibers with the skin using prestress, the device forms a stable electrical interface while preserving breathability and water resistance (Fig. 5b). The breathability of the fiber electrodes was evaluated by applying them to the forearm, as shown in Fig. 5c. The photos showed the air permeability of the bioelectronic fiber, which significantly reduces moisture accumulation and enhances comfort during prolonged wear. In contrast, traditional gel electrodes lack breathability, leading to discomfort, skin redness, and irritation, particularly after extended use. These characteristics are critical for long-term wear, effectively reducing discomfort caused by moisture accumulation and ensuring reliable signal transmission. Figure 5d compares ECG signals recorded using fiber electrodes and traditional electrodes. Due to the decreased electrode adhesion caused by motion and sweating, the traditional electrodes (black curve) exhibit noticeable baseline drift and noise interference. In contrast, the fiber electrodes (red curve) provide clear and stable ECG signals, capturing main physiological features such as the P wave, QRS complex, and T wave. Additionally, during motion, the fiber electrodes show excellent resistance to motion artifacts (Fig. 5e), maintaining signal quality even during repetitive movements37. This highlights their superior mechanical stability and noise suppression capabilities. Real-time monitoring experiments demonstrated the multifunctional fiber’s ability to record ECG signals before and after exercise continuously (Fig. 5f), clearly reflecting physiological changes caused by physical activity and showcasing its potential for dynamic physiological monitoring applications.

Beyond ECG monitoring, the multifunctional fiber was also used to acquire EMG signals for muscle activity monitoring. As illustrated in Figs. 5g and S17, electrical signals travel through neural pathways during muscle contraction to stimulate muscle fibers, causing potential changes inside and outside the muscle cells. These bioelectrical signals propagate to the skin surface and are detected by the electrodes of multifunctional fibers. Compared with traditional Ag/AgCl gel electrodes (Fig. 5h), the fiber electrodes demonstrated superior comfort and electrical contact stability, enhancing their reliability for long-term monitoring. Under muscle contraction, the fiber electrodes recorded clear and stable EMG signals (Fig. 5i) with consistent waveforms and amplitudes. The signals maintained high quality under both dry and wet conditions. Signal-to-noise ratio (SNR) analysis results (Fig. 5j) showed that the fiber electrodes retain a high SNR even in humid environments, significantly outperforming traditional electrodes. This validates their stability and anti-interference capabilities under complex conditions.

Additionally, the multifunctional fiber effectively captured EMG signals corresponding to muscle contractions of varying forces (Fig. 5k), with signal amplitude increasing proportionally to the contraction force. A hand gesture recognition experiment was conducted further to validate the multifunctional fiber’s multi-channel monitoring capabilities. As shown in Fig. 5l, the fiber electrodes successfully recorded four-channel EMG signals corresponding to six different hand gestures hand opening, clenching, and bending in the forward, backward, leftward, and rightward directions. The signals from each channel were clear, independent, and non-interfering, accurately reflecting the muscle activity characteristics of each gesture. The collected data were further processed using classification algorithms, generating a confusion matrix (Fig. 5m). The results showed high classification accuracy, with several gestures achieving 99.6% average recognition accuracy, confirming the potential of fiber electronic devices for high-precision gesture recognition and motion classification.

The multifunctional fibers were applied in vivo nerve stimulation experiments to validate their stable interfacing with biological tissues and precise control of muscle activity. As shown in Fig. 6a, the schematic illustrates the placement of fiber electrodes on the rat’s sciatic nerve. Due to its stretchability and softness, the thin multifunctional fiber can be wrapped around the sciatic nerve. The conductive pathways within the fibers ensure efficient charge transport, while the PEDOT: PSS-coated electrode surface further reduces impedance and optimizes the electrical contact interface with the nerve tissue. The fiber electrodes transmit electrical stimulation to the nerve by effectively converting voltage pulses into current, inducing muscle contraction. The electrodes were surgically positioned on the sciatic nerve and securely bound, forming a stable and tightly adhered interface (Fig. 6b). This robust connection ensures a stable electric field in the nerve region, enabling controlled stimulation and modulation of neural activity.

The nerve stimulation was visualized by monitoring the rat’s hindlimb movement (Fig. 6c). In the initial state (left image), the hindlimb remains relaxed without movement (Supplementary Video 4). The hindlimb undergoes a significant inward contraction upon applying electrical stimulation through the fiber electrodes (middle image). To evaluate the dynamic response of hindlimb movement, continuous tracking was conducted under different stimulation frequencies (Fig. 6d). At low frequencies, such as 1 Hz and 5 Hz, the hindlimb exhibits stable and periodic flexion and extension with consistent amplitude and clear waveforms, indicating the high reliability and control precision of the fiber electrodes. Figure 6e further quantifies the maximum movement angle under varying stimulation frequencies (200 mV and 500 mV voltage pulses). At low frequencies (e.g., 1 Hz), the hindlimb initially extends outward (extension), but as the frequency increases, the movement gradually transitions to flexion (inward bending). At higher frequencies (above 50 Hz), the movement amplitude significantly decreases, reflecting the physiological limit of the nerve-muscle system in response to high-frequency stimulation. Compared to 200 mV, the 500 mV condition produced greater movement amplitudes, validating the enhanced nerve activation efficacy at higher voltages. Figure 6f further quantifies the stimulation success rate using a 500 mV voltage pulse across a 1–100 Hz frequency range. The results show that at low frequencies (1–10 Hz), the success rate approaches 100%. However, as the frequency increases, the success rate gradually decreases, with a notable drop observed at 30 Hz, highlighting muscle fatigue and the diminishing response of the system under rapid stimulation. Additionally, the relationship between stimulation voltage and hindlimb movement response was explored (Fig. 6g). As the voltage increased from 50 mV to 1000 mV, the movement angle significantly increased, and the success rate improved accordingly, reaching nearly 100% at voltages above 200 mV. This result confirms that the fiber electrodes can trigger effective neural responses at relatively low voltages, with enhanced stimulation performance at higher voltages.

To evaluate the stability of the fiber electronic devices in physiological conditions, they were soaked in PBS solution for 2 days, and their stimulation performance was compared with the initial state (Fig. 6h). The results show no significant change in either maximum movement angle or success rate at 200 mV and 500 mV after soaking, demonstrating the excellent durability and stability of the fiber electrodes in simulated biological environments. The biocompatibility of the fiber electronic devices was further validated through cell culture assays using HEK293T cells. Figure 6i shows fluorescence images of cells cultured in the PBS control group, the conductive fiber group, and PEDOT-coated electrodes. Cells exhibited good viability in all groups, with intact morphology and uniform distribution, particularly in the conductive fiber and PBS groups (Fig. S18). Quantitative analysis of fluorescence intensity (Fig. 6j) revealed a slight decrease in the MWCNT electrode group, which was significantly improved after PEDOT:PSS modification. Ultimately, the fluorescence intensity in both the conductive fiber and electrode groups showed no significant difference compared with the PBS control group. These results demonstrate the excellent biocompatibility of the fiber electronic devices, with minimal impact on cell viability, making them well-suited for long-term biomedical applications.

Methods
Materials
The chemical reagents and elastomers used in this study were commercially sourced. These included thermoplastic polyurethane (TPU, Tecothane Soft AR-62A) form Lubrizol Inc. hydrogenated styrene-ethylene-butylene-styrene (SEBS, Tuftec H1221) from Asahi Kasei, and Ecoflex 00-30 from Smooth-On Inc. The chemical reagents comprised 3,4-ethylenedioxythiophene (EDOT), poly(4-styrenesulfonate sodium) (NaPSS), phosphate-buffered saline (PBS, 0.01M, pH 7.4), copper sulfate pentahydrate, disodium ethylenediaminetetraacetic acid (EDTA), triethanolamine, sulfuric acid, Triton X-100, and dimethylamine borane, all obtained from Shanghai Aladdin Biochemical Technology Co., Ltd. Multi-walled carbon nanotubes (MWCNTs, XFQ041) with a diameter of 10–20 nm were purchased from Nanjing XF NANO Materials Tech Co., Ltd. Galinstan was prepared by melting a metal mixture of Ga, In, and Sn in a weight ratio of 68.5:21.5:10 in a glove box at 80 °C for 2 h.

Fabrication of the conductive fiber electronics
The TPU fibers were fabricated using a desktop melt spinning system (model: HZ-RF-01) from Huizhi Electrospinning Co., Ltd. TPU pellets were melted in an extruder at 220 °C and then stretched and cooled at a speed of 20 m/min using a draw mechanism. Finally, the fibers were wound and collected using a winder. To deposit PDA on the fiber surface, the TPU fibers were immersed in a dopamine hydrochloride (2 mg/mL) solution in 50 mM Tris buffer (pH 8.5) and shaken continuously for 14 h. After deposition, the fibers were rinsed with deionized water to remove excess PDA and subsequently immersed in an aqueous silver nitrate solution (0.2 M) for 3 h, facilitating the growth of AgNPs as a seed layer. The fibers were then immersed in a chemical bath for copper deposition. The bath solution contained 0.0258 M copper sulfate pentahydrate, 0.0258 M disodium ethylenediaminetetraacetic acid (EDTA), 0.268 M triethanolamine, 0.09 M sulfuric acid, 0.1 w/w% Triton X-100, and 0.074 M dimethylamine borane. After 30 min of reaction, ~1 µm thick copper film was deposited in regions containing the AgNP seed layer. To create a liquid metal coating, bulk Galinstan was roll-coated over the copper film on the fiber while the substrate was immersed in a dilute hydrochloric acid solution (4 w/v%). During the alloying reaction, the copper film gradually dissolved into the liquid metal, forming a compliant, liquid-phase conductive coating on the fiber.

Fabrication of the fiber electrodes
The preparation process of the MWCNT/SEBS nanocomposite involved dispersing 50 mg of MWCNT into 10 mL of toluene using ultrasonication for 20 min, followed by thorough mixing with 1.5 g of SEBS solution [20% (w/w) in toluene]. The uncured MWCNT/SEBS nanodispersion was then fed into the liquid nozzle of a spraying device using a syringe pump, while 0.2 MPa compressed air was supplied to the gas nozzle to atomize the liquid into fine droplets. This process enabled the selective deposition of a smooth and uniform MWCNT/SEBS composite coating on the unmasked fibers. The coated nanocomposite was placed in an oven and heated at 70 °C to ensure complete drying. Subsequently, PEDOT:PSS was electrochemically deposited onto the MWCNT/SEBS nanocomposite electrodes to form an interfacial layer. The electrodeposition was performed using a three-electrode setup with an electrochemical workstation (Interface 1010E). The individual MWCNT/SEBS nanocomposite electrode served as the working electrode, a graphite rod as the counter electrode, and an Ag/AgCl electrode as the reference. An aqueous electrolyte solution containing 0.01 M EDOT and 0.1 M PSS was used for cyclic voltammetry electrodeposition. The deposition process involved scanning the potential from −0.5 V to 1.2 V at a rate of 50 mV/s for 30 cycles.

Characterization of fiber electronics
The microstructure of the fibers was examined using an optical microscope and a scanning electron microscope (SEM, Thermo Fisher Apreo 2S SEM). Mechanical properties were evaluated using an Instron universal testing machine (Model 5944). All tensile measurements were conducted at a pulling rate of 10 mm/min using fiber samples with a length of 5 cm. Electrochemical characterization was performed with an Interface 1010E potentiostat in a three-electrode configuration. The electrochemical stability window was measured in PBS solution using linear sweep voltammetry at a scan rate of 5 mV/s. Electrochemical impedance spectroscopy (EIS) was conducted at open circuit potential with an amplitude of 5 mV over a frequency range of 0.1 Hz to 1 MHz.

Epidermal electrophysiological signal recording
Informed and written consent was obtained prior to participation in the experiments by all subjects. The research process was harmless to humans. The phenomena measured represent natural physiological responses and are not influenced by sex. Because sex was not included as a variable in the study design, the results are not specific to any particular sex. Fiber electronics were connected to an Intan RHD2132 amplifier (Intan Technologies LLC) to record ECG and EMG signals. All channels were band-pass filtered, pre-amplified, and recorded at a sampling rate of 10 kHz. Before recording, the skin was cleaned with alcohol. For long-term and comparative tests, Red Dot 2223 electrodes (3M, USA) with a surface contact area of 2.4 cm² were used as the control group. Single-channel electrodes employed bipolar recording to capture electrophysiological signals. Conductive gel was applied to all electrode-skin interfaces to enhance signal quality. The wearable-related experiments were conducted under the guidance of Peking University People’s Hospital and were approved by the Institutional Review Board of Peking University People’s Hospital (approval no. 20240118).

In vivo animal experiments
All animal protocols were approved by the Tab of Animal Experimental Ethical Inspection of the Sir Run Run Shaw Hospital, Zhejiang University School of Medicine and conformed to the Guidelines for the Care and Use of Laboratory Animals published by the National Institutes of Health. The Number of Animal Use Permit was SRRSH2025-0018. A total of two adult male C57BL/6J mice (8 weeks old, 20 g) were used in this study. Since the experiment was solely a proof-of-concept demonstration of the fiber electrode for neural electrical stimulation, sex was not a relevant factor and therefore was not considered. Mice were anesthetized with isoflurane, and a 37 °C heating pad was placed beneath them. A 2 cm incision was made on the skin, and the muscles near the femur were carefully separated to expose the sciatic nerve. Fiber electrodes or platinum wires were bound to the exposed sciatic nerve for electrical stimulation. For hindlimb movement experiments, biphasic charge-balanced rectangular voltage pulses were applied to the sciatic nerve using a function generator (UTG900E, Uni-Trend Technology Co., Ltd.). Various voltage amplitudes (50 mV to 500 mV) were used to stimulate the nerve, and toe movement trajectories were recorded simultaneously. The success rate of toe movements was calculated as the percentage of successful toe responses relative to the total number of stimulation pulses. Success rates were measured across different stimulation frequencies. After the experiments, the mice were euthanized immediately.

Biocompatibility test
The biocompatibility of the fiber electronics was evaluated using 293T cells through cell viability (CCK-8) and live/dead staining assays, with material extract solutions prepared at a concentration of 20 mg/mL. The extract solution was prepared by incubating the material in a culture medium at 37 °C for 24 h, followed by filtration through a 0.22 µm membrane to remove particulates. Cells were seeded into a 96-well plate in standard culture medium and incubated for 24 h to allow attachment. The culture medium was then replaced with the prepared extract solution, and the cells were further incubated for 24 h. For the CCK-8 assay, 10 µL of CCK-8 reagent was added to each well after incubation, and the plate was incubated for an additional 3 h at 37 °C. Cell viability was quantified by measuring the optical density (OD) at 450 nm using a microplate reader. For live/dead staining, the culture medium was removed, and the cells were washed with PBS before adding a live/dead dye solution. The cells were incubated with the dye for 30 min at 37 °C, and fluorescence signals were observed using a confocal microscope, where live cells stained green and dead cells stained red. These methods provided a comprehensive assessment of the material’s biocompatibility over the 24-h exposure period.

Peer review rounds
[{"round": 1, "reviewer_comments": {"reviewer_1": "This manuscript presents a solution deposition fabrication strategy for the production of stretchable 1D liquid metal fibers. The authors also demonstrated its applications in fields such as electrophysiological signal sensing, in vivo neural stimulation, and wireless energy transmission. This is an interesting work for the development of soft electronics, but the following issues need to be addressed:\n1. The author claims that the wettability of liquid metals can be improved through alloying. Since silver can alloy with liquid metals, why is copper deposited on silver?\n2. Why is the stretching pattern of the red curve in Figure 2i different from the others?\n3. How is the stability of the prepared fibers under multiple cyclic (such as 1000 cycles) tensile tests? Only less than 100 cycles are shown in Figure 3f.\n4. The author claims that the prepared fibers can be integrated with textiles. Can the integrated fabric still maintain its original function after washing?\n5. Many studies on conductive fibers (including liquid metal fibers) have been reported. What are the advantages of the fibers prepared by the author compared to them?\n6. In Figure 5, the author claims that the prepared fiber electronic can suppress signal noise. I am curious about how the author stabilizes the electrical transmission between electronic and the skin, while the prepared electrodes seem to have no adhesion.\n7. How did the author construct a four-channel electromyography electrode on fibers (Figure 5l)? Relevant schematic diagrams need to be supplemented.\n8. The author claims in the introduction that fibers with a diameter of 50 microns can be prepared, but it is not shown in the article. However, according to the author's description, the packaging thickness is 25 microns, which increases the diameter of the packaging to 50 microns. Is the manuscript contradictory?\nOther minor issues:\n1. What is the scale bar length in Figure 1 and Figure 3?\n2. The caption for Figure 3i is missing.\n3. Some formatting issues need to be corrected.", "reviewer_2": "This study presents a method for achieving liquid metal (LM) adhesion on TPU fibers by chemically depositing copper on the surface of polymer fibers, followed by leveraging the wetting behavior between gallium and copper. Based on this principle, the authors fabricate a flexible conductor with multiple potential applications, including physiological signal monitoring, in vivo neural stimulation, and wireless energy transfer. However, the manuscript contains many critical issues that require comprehensive clarification and detailed experimental elaboration. Therefore, the current version of the manuscript falls short of the standards for publication and necessitates major revisions to address the significant methodological and interpretational shortcomings. Therefore, I recommend a rejection for this version. However, a resubmission might be reconsidered if the author could completely re-design their experiments and re-write the manuscript.\n1. Numerous high-quality studies on fiber-based sensors have already been published in the literature (the references listed below). However, the current manuscript, along with its supporting information, lacks a thorough and systematic comparison with these existing works in terms of key performance metrics (such as gauge factor, detection limit, and sensing range). The absence of such comparative analysis and quantitative benchmarking makes it difficult to assess the claimed advantages of the proposed sensor. Therefore, the authors are strongly advised to provide relevant performance data and engage in a more rigorous comparison to substantiate the novelty and superiority of their work.\n(1) Advanced Materials 36.11 (2024): 2304876.\n(2) Advanced Materials 36.32 (2024): 2404506.\n(3) Science advances 11.12 (2025): eadt0318.\n2. The processes of electroless copper deposition (e.g. Science 262.5140 (1993): 1711-1712. and Electrochimica Acta, 2019, 307: 536-542) and the wetting behavior of copper with gallium (e.g. Nature communications 13.1 (2022): 4763.) have been extensively studied and well-documented in prior literature. Therefore, the fabrication mechanism employed in this work is not novel and cannot be considered as the authors’ original contribution. The authors should clearly explicitly clarify how their approach differs from or advances beyond established methods.\n(3) In Figure 2b, the deposition of Galinstan (liquid metal) on the fiber surface following electroless copper plating is achieved by stretching the fiber through a Galinstan droplet in an HCl solution. However, it remains unclear whether key parameters—such as the stretching rate, the mechanical modulus of the fiber, the ambient temperature, and the volume or mass of the Galinstan droplet—affect the quality, uniformity, or adhesion of the resulting alloy coating. The authors should perform additional experiments and provide corresponding data analysis to clarify how these factors influence the deposition process and the performance of the final structure.\n(4) In the strain–resistance response curve presented in Figure 3f, the gauge factor (GF), a key indicator of sensor sensitivity, is not explicitly provided. While the use of R/R0 is common for describing resistance changes in flexible conductors, ΔR/R0 is more appropriate for evaluating strain sensors. Based on the data in Figure 3f, the GF of the presented sensor is approximately 4, which suggests relatively low sensitivity. Furthermore, the detection limit of the sensor—a critical performance metric—has not been reported. The omission of these fundamental parameters significantly hinders a comprehensive evaluation of the sensor’s performance.\n(5) In the inset of Figure 3f, the fiber sensor is subjected to only 80 stretching cycles, which is insufficient to rigorously evaluate its durability or long-term mechanical stability. Moreover, the applied strain value during these cycles is not indicated, making it difficult to assess the test conditions. This raises concerns about the overall reliability and robustness of the sensor. In addition, the caption for Figure 3i appears to be missing from the figure legend, which further compromises clarity and completeness of data presentation.\n(6) In Figure 2d, SEM mapping is presented for the fiber surface; however, the authors should also provide the specific elemental composition ratios of the deposited metals (Ag, Ga, In, Sn), along with the corresponding EDS spectra to support the elemental analysis. Furthermore, the adhesion between the metallic coating and the fiber surface is a critical factor for practical applications. The manuscript currently lacks any experimental characterization of the interfacial adhesion. Can the authors confirm whether the metal layer is prone to delamination under mechanical stress? It is strongly recommended that cross-sectional SEM images of the fiber sensor be included to directly visualize the metal–fiber interface and assess coating adhesion. Additionally, a supplementary video demonstrating that the Ga layer does not detach during stretching would greatly enhance the credibility of the mechanical stability claims.\n(7) The authors introduced a liquid metal (LM, Galinstan) layer onto the surface of Cu-plated fibers, which effectively reduced the surface resistance, as shown in Figures 3g – 3h. However, a lower initial resistance does not directly correlate with enhanced sensitivity in strain sensors. Given the relatively low gauge factor (GF = 4, Figure 3f) of the fabricated fiber, the device functions more as a strain-insensitive flexible conductor than a highly sensitive sensor. It remains unclear whether omitting the LM coating could potentially enhance the sensor’s sensitivity. Therefore, the authors are strongly encouraged to perform comparative experiments with and without LM deposition to clarify the impact of LM on sensing performance.\n(8) In Figure 5, the authors claim that the flexible fiber can detect and distinguish various weak external stimuli (e.g., pulse signals and hand gestures). However, based on the low gauge factor (GF = 4) shown in Figure 3f and the data presented in Figure 5b, it remains unclear whether the observed resistance responses originate from the fiber itself or from the MWCNT/SEBS–PEDOT:PSS conductive electrode applied on its surface. Given the high electrical conductivity of the LM-based fiber, it is plausible that it primarily functions as a conductive electrode, while the MWCNT/SEBS–PEDOT:PSS realizes the sensing functionality. This raises a significant ambiguity regarding the actual sensing component in the system. The authors should conduct proper control experiments to rule out this possibility and clearly identify the source of the sensing response.", "reviewer_3": "In this manuscript, the authors have demonstrated a multifunctional bioelectronic fiber fabricated by employing an elastomeric core fiber and subsequent layer-by-layer chemical solution deposition processes. The developed fiber electrode exhibits excellent mechanical stretchability, high electrical conductivity, and comprehensive potential for utilization in wearable and implantable bioelectronic applications.\nNonetheless, there remain several critical issues and experimental details that should be addressed clearly to make this manuscript suitable for publication in a high-impact journal such as Nature Communications. Therefore, a major revision is recommended.\n1. The coating of liquid metals onto yarns with polymeric or metallic surfaces has already been reported in previous studies. (https://doi.org/10.1021/acsnano.1c05762, https://doi.org/10.1002/adfm.202105587, https://doi.org/10.1021/acsami.9b23083). Therefore, it is essential to clarify and emphasize the distinctive novelty of your approach compared to these prior studies. This differentiation should be explicitly highlighted.\n2. In the fabrication process of the fiber electrodes, multiple layers (e.g., PDA, AgNP, and Cu) are sequentially deposited. Each layer is expected to significantly influence the mechanical and electrical properties of the fibers. Detailed characterization of changes in fiber diameter, stretchability, and electrical conductivity at each fabrication stage should be provided.\n3. The thickness and deposition amount of the internal layers (PDA, AgNW, and Cu layer) could significantly impact the quantity and uniformity of the liquid metal coating. Is there any relevant data available?\n4. The title and introduction emphasized \"scalable fabrication,\" which commonly implies suitability for mass production. However, the manuscript currently lacks experimental results whether fibers exceeding several meters can actually be fabricated, nor whether the diameter uniformity and electrical properties of the fibers can be maintained along their length. If such data cannot be provided, the use of the term \"scalable\" is not appropriate.\n5. The authors introduced additional functional layers (SEBS/CNT and PEDOT:PSS) onto the fiber electrodes. However, the electrical and electrochemical properties (Fig. 3e and Fig. 3g) are strongly dependent on the amount of PEDOT:PSS. The amount or weight% of the PEDOT:PSS should be clearly specified.\n6. The authors demonstrated several interesting applications, such as textile-based wireless power transmission (inductor coil), ECG, and EMG signal monitoring. Specifically, the electrodes shown in Fig. 5a and Fig. 5g are designed with 2-plied fibers. Was there any signal interference (crosstalk) between the two adjacent fiber electrodes?\n7. The authors also have presented an intriguing application involving PEDOT:PSS-coated electrodes for neural stimulation. However, PEDOT:PSS is amphiphilic and known to have potential leaching risks in wet physiological environments. Therefore, the authors should evaluate the resistance stability and amount of PEDOT:PSS under prolonged immersion conditions.\n8. When implanting electrodes in biological tissues, mechanical properties (stiffness or Young’s modulus) and weight of electrode are critically important because overly stiff or heavy electrodes can damage surrounding tissue. However, these critical mechanical parameters have not been sufficiently addressed in this manuscript.\n9. Overall, the manuscript is well-written. However, minor typographical errors remain (e.g., insperion, form Covestro, etc)."}, "author_response": "Point-by-point responses to the reviewers’ comments\nDear Reviewers,\nThanks for your consideration for publication of our work entitled \"Scalable and Stretchable 1D Bioelectronic Fibers for Multimodal Sensing and Stimulation\" (Manuscript ID: NCOMMS-25-48993-T).\nWe are very grateful for the detailed and thoughtful comments made by the reviewers. The comments are very helpful for us to further improve our work. The revision has been made to fully address the reviewer’s comments point by point. All revisions have been highlighted in red color in the revised manuscript for the convenience of reviewing.\nWe thank you in advance for your time and attention.\nWith Warm Regards,\nYours sincerely,\nYunlei Zhou\nReviewer #1\nThis manuscript presents a solution deposition fabrication strategy for the production of stretchable 1D liquid metal fibers. The authors also demonstrated its applications in fields such as electrophysiological signal sensing, in vivo neural stimulation, and wireless energy transmission. This is an interesting work for the development of soft electronics, but the following issues need to be addressed:\nWe thank the reviewer’s positive assessments and constructive suggestions. The detailed response to each point is provided in the following.\n1. The author claims that the wettability of liquid metals can be improved through alloying. Since silver can alloy with liquid metals, why is copper deposited on silver?\nResponse: We fully agree with the reviewer’s insightful comment. Indeed, silver is known to react with liquid metals to form Ag₂Ga. However, previous studies have reported that silver, even at the nanoscale, is not strongly infiltrated by liquid metals, which differs from the more complete and robust interaction observed between copper and liquid metals (Adv. Mater. Interfaces 2021, 8, 2100884; Journal of Applied Physics, 2005, 98(7)). In our work, silver ions are first reduced by the catechol groups of PDA to form silver nanoparticles, which act as catalytic nucleation sites for the subsequent electroless deposition of copper films. Compared to silver nanoparticles, continuous copper films can more effectively interact with liquid metals, enabling the formation of a more uniform liquid-metal coating. This observation is further confirmed by our experiments: silver nanoparticles generated solely via chemical reduction are insufficient to promote the uniform formation of liquid-metal fibers around the fibers, as shown in Supplementary Figure S1 of the revised manuscript.\nIn the revised supplementary information,\nFigure S1. The Ag particles in the fibers were unable to react with the liquid metal, and the fiber surface remained free of liquid metal after exposure.\n2. Why is the stretching pattern of the red curve in Figure 2i different from the others?\nResponse: We thank the reviewer for pointing out the issue with the mechanical cycling curves. Upon careful examination of the initial experimental data and records, we believe that the observed inconsistency likely resulted from minor variations during early sample preparation or clamping, which caused slight deviations in the behavior of the specific sample under 100% strain. We fully agree with the reviewer that this affected the consistency of the trends presented in the figure. Following the reviewer’s suggestion, we conducted additional experiments. The samples were re-prepared, and multiple repeated tests were performed under identical conditions for all strain levels. The new results show a high degree of reproducibility, with all stress and strain curves, including those at 100% and 150% strain, now exhibiting excellent consistency. The updated data have replaced the original Figure 2i.\nIn the revised manuscript,\n3. How is the stability of the prepared fibers under multiple cyclic (such as 1000 cycles) tensile tests? Only less than 100 cycles are shown in Figure 3f.\nResponse: We thank the reviewer for the valuable comment. The initial assessment based on 80 cycles was suboptimal for evaluating stability. Following the suggestion, we performed extended cycling up to 1000 cycles, which confirmed satisfactory electrical stability of the device. This observed stability aligns well with the established understanding that such liquid-metal-based multilayer electrodes are inherently stable.\nIn the revised manuscript, “f, Normalized resistance as a function of applied strain, ranging from 0% to 50%.”\n4. The author claims that the prepared fibers can be integrated with textiles. Can the integrated fabric still maintain its original function after washing?\nResponse: We greatly appreciate the reviewer’s important comment. In response, we have conducted additional experiments, the results of which are provided in Supplementary Figure S16 of the revised supplementary information. As shown, the washing process has a negligible effect on the electrical performance of the fabric. This stability is mainly attributed to the presence of a relatively thick encapsulation layer.\nIn the revised supplementary information,\nFigure S16. The washing process has a negligible impact on the electrical properties of the fabric. (a, b) Images demonstrating the washing procedure. (c) The change in electrical resistance as a function of multiple washing cycles.\n5. Many studies on conductive fibers (including liquid metal fibers) have been reported. What are the advantages of the fibers prepared by the author compared to them?\nResponse: We thank the reviewer for this thoughtful question. One-dimensional fibers have garnered significant interest for wearable and implantable applications owing to their compact structure, which imparts high flexibility, isotropic deformability, breathability, and lightweight characteristics. However, existing conductive fibers based on nanomaterials or conventional metals often struggle to simultaneously achieve high conductivity and robust mechanical flexibility(Table S1). In recent years, liquid metals have emerged as a promising material class to address this challenge, thanks to their high electrical conductivity and inherent softness.\nDespite this promise, the development of liquid metal fibers still faces considerable challenges. Unlike large area thin films or textile-based assemblies, processing at the single fiber level involves inherent complexities. At the microscale, achieving continuous and uniform liquid metal coatings requires highly controlled electrochemical deposition processes. The high surface tension of liquid metals further complicates their direct adhesion to elastic polymer substrates. Although the formation of a gallium oxide layer can moderately reduce surface tension, it offers limited precision for controlling interfacial properties at small scales, often resulting in nonuniform coatings. Alternative methods such as injection molding or thermal drawing are constrained by minimum fiber diameter limits typically not below 250 µm and often require encapsulation layers that prevent direct contact with biological tissues, thereby limiting sensing fidelity.\nIn summary, liquid metal fibers currently stand out as the only candidate capable of unifying electrical and mechanical properties. However, the predominant method for producing them is core-shell injection (Table S1), which insulates the conductive core and prevents direct contact for biosensing. Furthermore, coating methods lack the precision required for fine control.In this study, we present a scalable fabrication strategy for stretchable one-dimensional bioelectronic fibers. Using a layer-by-layer solution deposition process, we achieved uniform and adherent liquid metal coatings on ultrafine stretchable fibers, yielding high electrical conductivity and mechanical durability. The resulting fibers exhibit stable electrical performance with minimal resistance variation under various deformation modes including bending, twisting, and repeated stretching demonstrating strong potential for use in wearable electronics. Furthermore, our approach extends beyond the fabrication of individual conductive fibers. We have seamlessly integrated inert electrodes to construct a liquid metal fiber based conductive sensing system capable of centralized signal acquisition from single to multiple fibers. This platform not only enables the scalable production of high-performance LM fibers but also establishes a versatile foundation for advanced applications.\nTo address these points, we have made amendments to the corresponding parts of the main text.\nIn the revised manuscript, “Although various solution-based methods for preparing liquid metals have been reported, the large-scale fabrication of single liquid metal fiber remains unexplored37-39. The main challenge arises from the structural differences between films and fibers: large area films possess abundant reactive interfaces that readily interact with liquid metals, whereas single fibers have a much smaller surface area and thus require higher quality electrochemical coatings to achieve uniform and stable encapsulation. Moreover, individual liquid metal electrodes are highly electrochemically active, making them unsuitable for direct application in physiological environments. To address this issue, we introduced an inert CNT/PEDOT:PSS electrode layer, allowing the exposed regions to directly contact biofluids and perform sensing functions. This bilayer electrode structure, composed of liquid metal and inert materials, combines excellent conductivity with stable sensing performance in complex biological conditions. Furthermore, by twisting and assembling multiple encapsulated liquid-metal electrodes, a multichannel sensing system can be constructed (Fig. 1b).”\nIn the revised supplementary information,\nTable S1. Summary of recent studies on single conductive fibers\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nMelt Wearable\nspinning heater\nversatile\nintelligent\nwearable\napplications\nMelt\nspinning\nElectromagneti\nCNT/ Melt c Interference\nMXene spinning Shielding\nFabrics\nPEDOT:PSS Melt\n5 - - 3828 15 36 435\nspinning\nstretchable\nMXene/CN Melt\nT spinning\nnanogenerator\nGas sensor,\nCotto Melt\nn spinning\nsensor\nhierarchica\nElectrode,\nAgNWs/MX via lly\nene over-twisting interlocke\nsensor\nd helical\nJoule heaters,\nElectrothermoc\nPVD Coaxial\nF injected\ndisplays, and\nself-powered\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nwearable\nsensors\nOuter\nTriboelectric\nnanogenerator\nbonding\nElectrodes,\ncommunication\nfirefighting\nsuits\nStrain sensor,\nelectrode\nPDM Mold\nS injected\nOuter Physiological\nbonding Smart clothes\nStrain sensor,\nPressure\nCoaxial\ninjected\nswitches, Soft\nheater\nElectric heater,\nStrain sensing,\nCoaxial Pressure\ninjected sensing,\nProximity\nsensing\nCapillary Sustainable\ncoating\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nStrain sensing\napplications\nCoaxial electromagnetic\ninjected shielding\nCoaxial Physiological\ninjected Monitoring\nInductive\nEcofl Mold\nex injected\nNFC and\nMold\ninjected\nHarvesting\nStretchy Smart\nFabrics and\nCoaxial Electromagneti\ninjected c Interference\nShielding\nThi Multimodal\nFull solution Outer\ns Sensing and\nEGaIn TPU chemical layer 39623 50-100 250 1.85\nWo Stimulation\ndeposition bonding\nrk\n6. In Figure 5, the author claims that the prepared fiber electronic can suppress signal noise. I am curious about how the author stabilizes the electrical transmission between electronic and the skin, while the prepared electrodes seem to have no adhesion.\nResponse: Thank you for this insightful question. We encountered the same challenges during our experiments and have addressed them through two methods, as detailed in the manuscript. First, we applied a pre-tension to the elastic fibers to ensure stable contact with the skin. Second, a small amount of gel was applied to the interface to further enhance contact stability. As demonstrated in Figure 3h, these two approaches together enabled us to acquire stable epidermal electrical signals.\n7. How did the author construct a four-channel electromyography electrode on fibers (Figure 5l)? Relevant schematic diagrams need to be supplemented.\nResponse: We thank the reviewer for this comment. The multi-channel measurement capability was achieved by intertwining multiple fibers into a single strand, with each fiber serving as an independent electrode. A schematic diagram of this design is provided in the revised Supplementary Information, Figure S15. Design of the multichannel measurement fiber.\n8. The author claims in the introduction that fibers with a diameter of 50 microns can be prepared, but it is not shown in the article. However, according to the author's description, the packaging thickness is 25 microns, which increases the diameter of the packaging to 50 microns. Is the manuscript contradictory?\nResponse: We sincerely thank you for this valuable comment. Your observation is highly insightful and has helped us identify an ambiguity in our original description that required clarification. We apologize for any confusion caused. Please allow us to clarify the point regarding fiber diameter: the statement in the Introduction that we are able to \"fabricate fibers with a diameter of approximately 50 μm\" refers specifically to the final diameter of the composite fiber electrode after the deposition of all functional layers (including PDA, Ag, Cu, and LM). Starting from a thermoplastic polyurethane (TPU) core fiber with an initial diameter of less than 50 μm, the sequential deposition process allows us to control the total diameter to around 50 μm. The diameter distribution for each layer is provided in the figure below. If an encapsulation layer is applied, the overall diameter would increase to approximately 100 μm. We have carefully revised the corresponding section in the manuscript to reflect this clarification and ensure accuracy.\nIn the revised manuscript, “achieving a fiber diameter as small as 50-100 µm”\nIn the revised supplementary information,\nFigure S2. Microscopic details of the fiber at each fabrication stage. (a) Untreated TPU fiber exhibiting a smooth surface. (b) Fiber after 14 h of polydopamine (PDA) deposition, showing the formation of fine particulate structures while retaining overall smooth morphology. (c) Fiber after 3 h of AgNO₃ treatment, revealing the presence of discrete silver nanoparticles. (d) Fiber after 30 min of electrodes copper plating, displaying a dense and conglomerated copper coating. (e) Fiber after liquid metal attachment under HCl conditions, illustrating the successful binding of LM to the activated surface.\nOther minor issues:\n1. What is the scale bar length in Figure 1 and Figure 3?\nResponse: We appreciate the reviewer's careful attention to detail. In response, we have added the missing scale bar to the figure as indicated\n2. The caption for Figure 3i is missing.\nResponse: We thank the reviewer for the nice reminder, and the missing caption has been added accordingly.\n3. Some formatting issues need to be corrected.\nResponse: We thank the reviewer for pointing this out. We have carefully checked the manuscript and corrected the formatting issues in the revised version.\nReviewer #2\nThis study presents a method for achieving liquid metal (LM) adhesion on TPU fibers by chemically depositing copper on the surface of polymer fibers, followed by leveraging the wetting behavior between gallium and copper. Based on this principle, the authors fabricate a flexible conductor with multiple potential applications, including physiological signal monitoring, in vivo neural stimulation, and wireless energy transfer. However, the manuscript contains many critical issues that require comprehensive clarification and detailed experimental elaboration. Therefore, the current version of the manuscript falls short of the standards for publication and necessitates major revisions to address the significant methodological and interpretational shortcomings. Therefore, I recommend a rejection for this version. However, a resubmission might be reconsidered if the author could completely re-design their experiments and re-write the manuscript.\nWe thank the reviewer’s constructive suggestions. The detailed response to each point is provided in the following.\n1. Numerous high-quality studies on fiber-based sensors have already been published in the literature (the references listed below). However, the current manuscript, along with its supporting information, lacks a thorough and systematic comparison with these existing works in terms of key performance metrics (such as gauge factor, detection limit, and sensing range). The absence of such comparative analysis and quantitative benchmarking makes it difficult to assess the claimed advantages of the proposed sensor. Therefore, the authors are strongly advised to provide relevant performance data and engage in a more rigorous comparison to substantiate the novelty and superiority of their work.\n(1) Advanced Materials 36.11 (2024): 2304876.\n(2) Advanced Materials 36.32 (2024): 2404506.\n(3) Science advances 11.12 (2025): eadt0318.\nResponse: We sincerely thank the reviewer for raising this important point regarding the growing body of literature on fiber-based sensors. This is indeed a valuable observation, and we appreciate the opportunity to further clarify the positioning of our work. In direct response to the comment, we have now carefully compiled and incorporated a detailed comparison table in the revised supplementary information, which systematically summarizes the types and key distinguishing features of representative fiber-shaped electronic devices reported in the field. Overall, the fiber we developed distinguishes itself through a favorable profile that combines conductivity, structural adaptability, and mechanical performance. We believe this addition helps to better contextualize our contribution within the existing research landscape.\nFurthermore, we are grateful to the reviewer for drawing our attention to the three relevant papers, which we have now cited in the revised manuscript. We have studied these suggested works thoroughly and with great interest. It was found that the present manuscript was substantially different from those three published papers.\nThe first paper (Advanced Materials 36.11 (2024): 2304876) reports an indoor photovoltaic fiber using a highly conductive core-sheath titanium/carbon nanotube fiber. We note that our work, in contrast, focuses on fabricating liquid metal fibers for applications in electromyography and neural electrical stimulation. From a materials perspective, the cited work employs carbon nanotube-based electrodes, whereas our study develops conductive liquid metal electrodes for use under large strain. In terms of application, the prior work demonstrates a photovoltaic fiber, while our research is dedicated to health monitoring and stimulation. Apart from the shared form factor of fiber-based devices, the two studies address differing scientific and application goals, making them substantially distinct.\nThe second paper (Advanced Materials 36.32, 2024: 2404506) presents an interesting in situ strategy for fabricating hierarchical hydrogel fibers that exhibit circadian rhythmicity and memory capabilities, with their output voltage oscillating over a 24-hour cycle. In contrast, our work is focused on the development of highly stretchable conductive fibers based on liquid metals, which are specifically designed for applications in health monitoring and implantable devices. These two studies are based on completely different material systems (hydrogel vs. liquid metal) and pursue distinct application goals (biomimetic oscillations vs. bio-electrical interfacing), representing two separate research paths within the field of functional fibers.\nThe third article (Science Advances 11.12, 2025: eadt0318) reports a self-powered wearable haptic technology that utilizes a TENG as an energy source to create a comfortable, electrotactile textile glove. It is important to note that this work achieves its functionality at the fabric level, which is fundamentally different from our research and the two previously mentioned studies, all of which focus on engineering functional devices at the single-fiber level. This distinction in the fundamental building block (fabric vs. single fiber) clearly distinguishes the material platforms, fabrication methodologies, and target applications.\nIn conclusion, our work is distinct from the three referenced studies in multiple key aspects, including the materials system and targeted applications. The central emphasis of our research is on the scalable production of uniform liquid metal fibers and their implementation in wearable and implantable technologies. Motivated by the reviewer’s suggestion, we have incorporated this comparative discussion to underscore the unique contributions of our manuscript within the revised introduction. Additionally, we summarize recent advances in the field of conductive single fibers.\nIn the revised manuscript, “Although textile electronics have achieved considerable success in wearable applications, single fibers offer superior characteristics for particular applications6-13. ”\n“Although various solution-based methods for preparing liquid metals have been reported, the large-scale fabrication of single liquid metal fiber remains unexplored37-39. The main challenge arises from the structural differences between films and fibers: large area films possess abundant reactive interfaces that readily interact with liquid metals, whereas single fibers have a much smaller surface area and thus require higher quality electrochemical coatings to achieve uniform and stable encapsulation. Moreover, individual liquid metal electrodes are highly electrochemically active, making them unsuitable for direct application in physiological environments. To address this issue, we introduced an inert CNT/PEDOT:PSS electrode layer, allowing the exposed regions to directly contact biofluids and perform sensing functions. This bilayer electrode structure, composed of liquid metal and inert materials, combines excellent conductivity with stable sensing performance in complex biological conditions. Furthermore, by twisting and assembling multiple encapsulated liquid-metal electrodes, a multichannel sensing system can be constructed (Fig. 1b). ”\nIn the revised supplementary information,\nTable S1. Summary of recent studies on single conductive fibers\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nMelt Wearable\nspinning heater\nversatile\nintelligent\nwearable\napplications\nMelt\nspinning\nElectromagneti\nCNT/ Melt c Interference\nMXene spinning Shielding\nFabrics\nPEDOT:PSS Melt\n5 - - 3828 15 36 435\nspinning\nstretchable\nMXene/CN Melt\nT spinning\nnanogenerator\nGas sensor,\nCotto Melt\nn spinning\nsensor\nhierarchica\nElectrode,\nAgNWs/MX via lly\nene over-twisting interlocke\nsensor\nd helical\nJoule heaters,\nElectrothermoc\nPVD Coaxial\nF injected\ndisplays, and\nself-powered\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nwearable\nsensors\nOuter\nTriboelectric\nnanogenerator\nbonding\nElectrodes,\ncommunication\nfirefighting\nsuits\nStrain sensor,\nelectrode\nPDM Mold\nS injected\nOuter Physiological\nbonding Smart clothes\nStrain sensor,\nPressure\nCoaxial\ninjected\nswitches, Soft\nheater\nElectric heater,\nStrain sensing,\nCoaxial Pressure\ninjected sensing,\nProximity\nsensing\nCapillary Sustainable\ncoating\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nStrain sensing\napplications\nCoaxial electromagnetic\ninjected shielding\nCoaxial Physiological\ninjected Monitoring\nInductive\nEcofl Mold\nex injected\nNFC and\nMold\ninjected\nHarvesting\nStretchy Smart\nFabrics and\nCoaxial Electromagneti\ninjected c Interference\nShielding\nThi Multimodal\nFull solution Outer\ns Sensing and\nEGaIn TPU chemical layer 39623 50-100 250 1.85\nWo Stimulation\ndeposition bonding\nrk\n2. The processes of electroless copper deposition (e.g. Science 262.5140 (1993): 1711-1712. and Electrochimica Acta, 2019, 307: 536-542) and the wetting behavior of copper with gallium (e.g. Nature communications 13.1 (2022): 4763.) have been extensively studied and well-documented in prior literature. Therefore, the fabrication mechanism employed in this work is not novel and cannot be considered as the authors’ original contribution. The authors should clearly explicitly clarify how their approach differs from or advances beyond established methods.\nResponse: We sincerely thank the reviewer for raising this important point. We have carefully reviewed the cited literature, which primarily focuses on metal fabrication on polymer substrates and patterning of liquid metals via metallic templates. The knowledge from these studies can indeed be conceptually transferred to the patterned fabrication of liquid metals. However, our work is situated within the distinct field of fiber electronics. Although liquid metals have been widely studied, a method for their liquid-phase fabrication at the single-fiber level has not been reported. Notably, the adaptation of thin-film electronic techniques to single-fiber devices is significant enough to warrant publication in high-impact journals, including those in the Nature family, highlighting a key opportunity for innovation that our work aims to address. For instance, single-fiber electroluminescent devices have been reported (Nature Photon. 9, 233–238, 2015), even though flexible electroluminescent devices have already been widely studied. Similarly, single-fiber batteries have recently been demonstrated (Nature 629, 86–91, 2024), despite the extensive prior development of flexible battery technologies.\nIn terms of specific innovations in fabrication, current liquid metal fiber production methods primarily rely on directly coating or injecting liquid metal into hollow or thermoplastic fibers to form LM-core structures (Liquid Metal Fibers, Adv. Fiber Mater. 4, 987–1004, 2022). These approaches often suffer from uneven coating, relatively large diameters, or poor skin contact of the conductive components. Our study addresses these challenges by introducing a liquid-phase processing strategy for the fabrication of single LM fibers. Unlike large-area thin films or woven fabrics, the processing of individual fibers is inherently more complex. At the microscale, highly uniform electrochemical deposition is required to achieve continuous and stable LM fiber formation. In the revised manuscript, we have further investigated the influence of various fabrication parameters on LM deposition and fiber quality. Moreover, beyond the chemical deposition of LM fibers, we have seamlessly integrated inert electrodes to develop an LM-fiber-based conductive sensing system. In other words, our approach not only enables the fabrication of individual LM fibers but also establishes a complete LM-based sensing platform. This allows for centralized sensing from single to multiple fibers, thereby broadening their applicability in electronic systems. In response to the reviewer’s comments, we have revised and expanded the manuscript to better emphasize the novelty and originality of our work.\nIn the revised manuscript, “Although various solution-based methods for preparing liquid metals have been reported, the large-scale fabrication of single liquid metal fiber remains unexplored37-39. The main challenge arises from the structural differences between films and fibers: large area films possess abundant reactive interfaces that readily interact with liquid metals, whereas single fibers have a much smaller surface area and thus require higher quality electrochemical coatings to achieve uniform and stable encapsulation. Moreover, individual liquid metal electrodes are highly electrochemically active, making them unsuitable for direct application in physiological environments. To address this issue, we introduced an inert CNT/PEDOT:PSS electrode layer, allowing the exposed regions to directly contact biofluids and perform sensing functions. This bilayer electrode structure, composed of liquid metal and inert materials, combines excellent conductivity with stable sensing performance in complex biological conditions. Furthermore, by twisting and assembling multiple encapsulated liquid-metal electrodes, a multichannel sensing system can be constructed (Fig. 1b). ”\n3. In Figure 2b, the deposition of Galinstan (liquid metal) on the fiber surface following electroless copper plating is achieved by stretching the fiber through a Galinstan droplet in an HCl solution. However, it remains unclear whether key parameters—such as the stretching rate, the mechanical modulus of the fiber, the ambient temperature, and the volume or mass of the Galinstan droplet—affect the quality, uniformity, or adhesion of the resulting alloy coating. The authors should perform additional experiments and provide corresponding data analysis to clarify how these factors influence the deposition process and the performance of the final structure.\nResponse: We sincerely thank the reviewer for raising this critical point regarding process parameters. Their insightful observation demonstrates a deep understanding of materials processing and directly addresses a core aspect of our methodology, the precise control of coating quality. The four parameters highlighted—drawing speed, fiber modulus, ambient temperature, and Galinstan droplet volume—form the foundation for achieving reproducible, high-performance liquid metal coatings. This comment inspired what we consider one of the most valuable aspects of our revision: a full-factorial parametric analysis that has substantially strengthened the scientific contribution of our work.\nGuided by the reviewer’s suggestion, we conducted a systematic experimental campaign to examine both the individual and interactive effects of each parameter on coating uniformity, microstructural quality, and interfacial adhesion. Several key insights emerged, refining our processing protocol. Most notably, drawing speed proved to be the dominant factor. At speeds approaching 20 cm/s, the liquid metal–copper reaction becomes kinetically limited, resulting in non-uniform coverage and defective interfaces. This finding enabled us to define a precise operational window for high-speed coating. In contrast, the Young’s modulus of the elastomeric fiber substrate had a surprisingly negligible effect across the tested range, suggesting broad material compatibility for our method. Similarly, the Galinstan droplet volume minimally affected final fiber morphology, provided complete surface wetting was achieved, simplifying process control. Variations within a standard ambient temperature range also had no measurable impact, confirming the robustness of the deposition process under typical laboratory conditions. In response to the comments, we have conducted additional experiments and provided detailed information on the fabrication process as suggested.\nIn the revised manuscript, “Moreover, the corresponding stress-strain behavior was not significantly affected (Fig. S6). In contrast, the electrical conductivity changed markedly at each fabrication stage, as shown in Fig. S7. The initial Ag layer failed to form a stable conductive network, resulting in negligible conductivity. The subsequent Cu coating achieved a low conductivity of 0.1133 S cm⁻¹. Finally, after reaction with the liquid metal, the conductivity dramatically increased to 39623 S cm⁻¹. The thickness and deposition amount of the internal layers (PDA, AgNW, and Cu) were found to be highly dependent on the deposition time. As summarized in Fig. S8, the optimized parameters are as follows: a 14-hour PDA deposition formed a uniform and continuous adhesive layer, providing an ideal foundation for subsequent metallization. A 3-hour AgNP deposition achieved sufficient surface coverage to act as an effective conductive activation layer, while a 30-minute copper electroplating produced a continuous and dense metallic film that served as an excellent conductive substrate for the subsequent Galinstan coating. Collectively, these optimized parameters ensure ideal, full-surface coverage of the final liquid metal layer.”\n“The uniformity and quality of the liquid metal coating are crucial for ensuring the reliable performance of fiber electrodes. This deposition process is influenced by several key processing parameters. To establish a robust and reproducible process window, we systematically investigated the effects of drawing speed, fiber modulus, ambient temperature, and the volume of Galinstan droplets on the final coating morphology. In Fig. S3, at speeds approaching 20 cm/s, the reaction between the liquid metal and copper becomes kinetically limited, resulting in non-uniform coverage and a defective interface. This finding allows us to define a precise operating window for high-speed coating. In contrast, the Young’s modulus of the elastic fiber substrate, ambient temperature, and the volume of liquid metal exhibited unexpectedly minimal influence within the tested ranges, indicating broad material compatibility of our approach.”\nIn the revised supplementary information,\nTable S1. Summary of recent studies on single conductive fibers\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nMelt Wearable\nspinning heater\nversatile\nintelligent\nwearable\napplications\nMelt\nspinning\nElectromagneti\nCNT/ Melt c Interference\nMXene spinning Shielding\nFabrics\nPEDOT:PSS Melt\n5 - - 3828 15 36 435\nspinning\nstretchable\nMXene/CN Melt\nT spinning\nnanogenerator\nGas sensor,\nCotto Melt\nn spinning\nsensor\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nene over-twisting lly humidity\ninterlocke sensor\nd helical\nJoule heaters,\nElectrothermoc\nhromic\nPVD Coaxial\nF injected\nself-powered\nwearable\nsensors\nOuter\nTriboelectric\nnanogenerator\nbonding\nElectrodes,\ncommunication\nfirefighting\nsuits\nStrain sensor,\nelectrode\nPDM Mold\nS injected\nOuter Physiological\nbonding Smart clothes\nStrain sensor,\nPressure\nCoaxial\ninjected\nswitches, Soft\nheater\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nElectric heater,\nStrain sensing,\nCoaxial Pressure\ninjected sensing,\nProximity\nsensing\nCapillary Sustainable\ncoating\nStrain sensing\napplications\nCoaxial electromagnetic\ninjected shielding\nCoaxial Physiological\ninjected Monitoring\nInductive\nEcofl Mold\nex injected\nNFC and\nMold\ninjected\nHarvesting\nStretchy Smart\nFabrics and\nCoaxial Electromagneti\ninjected c Interference\nShielding\nThi Multimodal\nFull solution Outer\ns Sensing and\nEGaIn TPU chemical layer 39623 50-100 250 1.85\nWo Stimulation\ndeposition bonding\nrk\nFigure S3. Schematic illustration of the key processing parameters investigated for the deposition process: (a) drawing speed, (b) fiber Young's modulus, (c) LM droplet mass, and (d) temperature.\nFigure S6. Stress-strain behavior of the fiber at different fabrication stages.\nFigure S7. Electrical conductivity of the fiber at different fabrication stages.\nFigure S8. Influence of deposition time on the morphological evolution of each functional layer. (a) A discontinuous PDA film formed after 3 h. (b) Discontinuous Ag nanoparticles appeared between 0- 2 h. (c) Corresponding elemental mapping for Ag. (d) A dense Cu layer formed within 30 min due to the fast deposition rate. (e) Corresponding elemental mapping for Cu.\n4. In the strain–resistance response curve presented in Figure 3f, the gauge factor (GF), a key indicator of sensor sensitivity, is not explicitly provided. While the use of R/R0 is common for describing resistance changes in flexible conductors, ΔR/R0 is more appropriate for evaluating strain sensors. Based on the data in Figure 3f, the GF of the presented sensor is approximately 4, which suggests relatively low sensitivity. Furthermore, the detection limit of the sensor—a critical performance metric—has not been reported. The omission of these fundamental parameters significantly hinders a comprehensive evaluation of the sensor’s performance.\nResponse: We sincerely appreciate the reviewer's insightful comments regarding the resistance variation rate of our liquid metal fibers. This valuable feedback allows us to better articulate the design rationale behind our work. We would like to clarify that the liquid metal network in our study was specifically designed to function as a stable conductive pathway, rather than serving as a strain-sensing element. To ensure precise signal acquisition, we integrated an inert electrode interface with this conductive network. This design approach is grounded in the distinct functional requirements of different components: strain sensors require significant resistance variations to achieve high sensitivity, while signal transmission electrodes demand minimal resistance changes to maintain signal fidelity. Therefore, the minimal resistance variation demonstrated by our liquid metal electrode under strain conditions actually represents a key advantage, confirming its ability to maintain stable electrical performance even under substantial deformation, which is a crucial characteristic for developing reliable bioelectronic interfaces. We are grateful for the opportunity to provide this clarification, which we believe helps to better demonstrate the design considerations underlying our work.\n(5) In the inset of Figure 3f, the fiber sensor is subjected to only 80 stretching cycles, which is insufficient to rigorously evaluate its durability or long-term mechanical stability. Moreover, the applied strain value during these cycles is not indicated, making it difficult to assess the test conditions. This raises concerns about the overall reliability and robustness of the sensor. In addition, the caption for Figure 3i appears to be missing from the figure legend, which further compromises clarity and completeness of data presentation.\nResponse: We sincerely thank the reviewer for this valuable suggestion. We agree that the initial stability assessment based on only 80 cycles was insufficient for a comprehensive evaluation. In response, we have conducted extended cycling tests up to 1,000 cycles under 0-50% strain conditions. The results confirm the device’s excellent electrical stability, consistent with the well-established robustness of liquid metal based multilayer electrodes in both structure and function.\nWe also appreciate the reviewer’s kind reminder regarding the missing caption, which has now been added in the revised manuscript. In addition, we have carefully checked the entire manuscript and corrected all identified errors to ensure clarity and accuracy.\nIn the revised manuscript, “f, Normalized resistance as a function of applied strain, ranging from 0% to 50%.”\n(6) In Figure 2d, SEM mapping is presented for the fiber surface; however, the authors should also provide the specific elemental composition ratios of the deposited metals (Ag, Ga, In, Sn), along with the corresponding EDS spectra to support the elemental analysis. Furthermore, the adhesion between the metallic coating and the fiber surface is a critical factor for practical applications. The manuscript currently lacks any experimental characterization of the interfacial adhesion. Can the authors confirm whether the metal layer is prone to delamination under mechanical stress? It is strongly recommended that cross-sectional SEM images of the fiber sensor be included to directly visualize the metal–fiber interface and assess coating adhesion. Additionally, a supplementary video demonstrating that the Ga layer does not detach during stretching would greatly enhance the credibility of the mechanical stability claims.\nResponse: We fully agree that quantitative elemental analysis is essential for confirming the chemical composition of the coating. In response, we have performed additional energy-dispersive X-ray spectroscopy line scanning analysis. The corresponding EDS spectrum and elemental mapping results have been included as new figures in the revised manuscript. The spectrum obtained from the selected area clearly displays the characteristic peaks of Ag, Ga, In, and Sn. Notably, Ga appears as the most abundant element, while a certain amount of Cu remains detectable after the formation of the Cu-Ga alloy. The data is included as Figure S6 in the revised version of supplementary information.\nIn the revised manuscript, “The Energy Dispersive X-ray Spectroscopy (EDS) spectrum acquired from the selected area clearly shows the characteristic peaks of Ag, Ga, In, and Sn. Ga is identified as the dominant phase, with the significant presence of Cu being consistent with the formation of a Cu-Ga intermetallic alloy(Figure S4).”\nIn the revised supplementary information,\nFigure S4. Quantitative elemental composition (Ag, Ga, In, Sn) of the liquid metal alloy with supporting EDS spectrum.\nWe have carefully prepared cross-sectional samples of the fiber and performed SEM imaging to directly observe the metal-fiber interface and evaluate coating adhesion. These results have been included as Figure S8 in the revised version of supplementary information. The magnified cross-sectional view clearly reveals a continuous and compact interface between the metal coating and the fiber substrate, with no visible gaps or cracks.\nIn the revised manuscript, “Furthermore, cross-sectional SEM imaging together with elemental analysis demonstrates the robust adhesion of liquid metal to the TPU fibers (Fig. S5).”\nIn the revised supplementary information,\nFigure S5. Cross-sectional characterization of the conductive fiber at higher magnification. (a) SEM image. (b) Elemental mapping of Ga.\nIn response to the reviewer's valuable suggestion regarding mechanical stability assessment, we have further produced a supplementary video recording (submitted as Supplementary Video SV1) to dynamically visualize the coating performance under mechanical stress. This video documentation captures the fiber sensor undergoing multiple cycles of stretching and bending deformation, during which the metallic coating maintains exceptional structural integrity without exhibiting any signs of cracking, delamination, or interfacial detachment. These observations provide compelling visual evidence of the coating's remarkable durability and adhesion strength under dynamic mechanical loading conditions. The combination of these static interfacial characterization and dynamic mechanical testing results offers a multi-faceted validation of both the structural quality and functional reliability of our metallic coating approach, substantially strengthening the experimental support for our methodology.\nIn the revised manuscript, “A complementary video (Supplementary Video SV4) was recorded to validate the superior durability and adhesion of our conductive fiber under dynamic mechanical stress. The footage shows the sensor undergoing repeated stretching and bending cycles. Notably, the metallic coating retained its structural integrity throughout, with no evidence of cracking, delamination, or interfacial detachment.”\nIn the revised Supplementary Video SV4,\n(6) The authors introduced a liquid metal (LM, Galinstan) layer onto the surface of Cu-plated fibers, which effectively reduced the surface resistance, as shown in Figures 3g – 3h. However, a lower initial resistance does not directly correlate with enhanced sensitivity in strain sensors. Given the relatively low gauge factor (GF = 4, Figure 3f) of the fabricated fiber, the device functions more as a strain-insensitive flexible conductor than a highly sensitive sensor. It remains unclear whether omitting the LM coating could potentially enhance the sensor’s sensitivity. Therefore, the authors are strongly encouraged to perform comparative experiments with and without LM deposition to clarify the impact of LM on sensing performance.\nResponse: We sincerely appreciate the reviewer's insightful comments, which have prompted us to elaborate more comprehensively on the research objectives. It is important to clarify that the primary goal of fabricating a liquid metal layer onto the surface of Cu-plated fibers was to develop a highly reliable liquid metal electrode capable of maintaining stable electrical conductivity under extreme deformation conditions, rather than to create a strain sensor. This design rationale is grounded in the unique material advantages of liquid metals: their inherent fluidity enables exceptional deformation tolerance, while their metallic nature ensures superior electrical conductivity. Experimental results confirm that the electrode maintains stable electrical performance even under large strain conditions. Unlike strain sensors, which operate based on resistance changes in response to mechanical deformation, our liquid metal electrode is designed specifically for reliable signal transmission. Signal transmission electrodes require minimal resistance variation to ensure signal integrity, whereas strain sensors rely on significant resistance changes for sensitivity detection. Guided by this design principle, we have successfully developed a fiber-shaped electrode that combines high conductivity with outstanding mechanical stability, offering an innovative solution for signal transmission in wearable devices. We believe this clarification helps to better articulate the innovations and potential applications of our work. We are grateful for the opportunity to provide this explanation, which we hope will more accurately convey the research intentions and contributions of our study.\n(7) In Figure 5, the authors claim that the flexible fiber can detect and distinguish various weak external stimuli (e.g., pulse signals and hand gestures). However, based on the low gauge factor (GF = 4) shown in Figure 3f and the data presented in Figure 5b, it remains unclear whether the observed resistance responses originate from the fiber itself or from the MWCNT/SEBS–PEDOT:PSS conductive electrode applied on its surface. Given the high electrical conductivity of the LM-based fiber, it is plausible that it primarily functions as a conductive electrode, while the MWCNT/SEBS–PEDOT:PSS realizes the sensing functionality. This raises a significant ambiguity regarding the actual sensing component in the system. The authors should conduct proper control experiments to rule out this possibility and clearly identify the source of the sensing response.\nResponse: We sincerely thank the reviewer for this insightful comment. We would like to respectfully clarify that the intention of our work was not to employ liquid etal electrodes as strain sensors for detecting or distinguishing various weak external stimuli. Rather, our study is centered on the broader application of bioelectrical signal acquisition and regulation. In particular, we utilize the excellent flexibility and conductivity of liquid metal electrodes to achieve high-fidelity recording of electromyography (EMG) and electrocardiography (ECG) signals. Furthermore, to realize stable in vivo electrical stimulation, we incorporate MWCNT/SEBS–PEDOT:PSS inert electrodes as a complementary component.\nThe core concept underlying our approach is to establish a robust and reliable signal transmission pathway. The liquid metal electrodes, owing to their unique fluidity and mechanical resilience, provide a stable electrical interface even under dynamic strain conditions. In parallel, in regions that may be susceptible to electrochemical corrosion or interfacial instability, the introduction of inert electrodes effectively addresses these challenges. This integrated strategy thus ensures the long-term stability, reliability, and functionality of the system.\nReviewer #3\nIn this manuscript, the authors have demonstrated a multifunctional bioelectronic fiber fabricated by employing an elastomeric core fiber and subsequent layer-by-layer chemical solution deposition processes. The developed fiber electrode exhibits excellent mechanical stretchability, high electrical conductivity, and comprehensive potential for utilization in wearable and implantable bioelectronic applications.\nNonetheless, there remain several critical issues and experimental details that should be addressed clearly to make this manuscript suitable for publication in a high-impact journal such as Nature Communications. Therefore, a major revision is recommended.\nWe thank the reviewer’s positive assessments and constructive suggestions. The detailed response to each point is provided in the following.\n1. The coating of liquid metals onto yarns with polymeric or metallic surfaces has already been reported in previous studies. (https://doi.org/10.1021/acsnano.1c05762, https://doi.org/10.1002/adfm.202105587, https://doi.org/10.1021/acsami.9b23083). Therefore, it is essential to clarify and emphasize the distinctive novelty of your approach compared to these prior studies. This differentiation should be explicitly highlighted.\nResponse: We sincerely thank the reviewer for raising this important point. We have carefully studied the suggested literature and recognize that these works indeed demonstrate the successful integration of liquid metal on fabric substrates for various applications. However, it is important to note that these systems achieve their functionality primarily at the fabric or textile level, whereas our research focuses specifically on engineering functional devices at the single-fiber level. This fundamental difference in the scale of the functional unit (fabric vs. single fiber) leads to distinctly different material platforms, fabrication methodologies, and target applications.\nDue to their superior flexibility, isotropic deformability, breathability, and lightweight nature, single fiber-based electronics are uniquely advantageous for integration with biological tissues. In this context, liquid metal fibers have attracted growing attention for their outstanding conductivity, intrinsic stretchability, facile phase transition, and the potential to be woven into smart textiles. Nevertheless, current liquid metal single fiber fabrication methods mainly rely on direct coating or injecting LM into hollow or thermoplastic fibers to form LM-core structures (Liquid Metal Fibers. Adv. Fiber Mater. 4, 987–1004 (2022)). These approaches often suffer from uneven coating, relatively large diameters, or poor skin contact of the conductive components. Our study addresses these challenges by introducing a liquid-phase processing strategy for the fabrication of single LM fibers. Unlike large-area thin films or woven fabrics, the processing of individual fibers is inherently more complex. At the microscale, highly uniform electrochemical deposition is required to achieve continuous and stable LM fiber formation. In the revised manuscript, we have further investigated the influence of various fabrication parameters on LM deposition and fiber quality. Moreover, beyond the chemical deposition of LM fibers, we have seamlessly integrated inert electrodes to develop an LM-fiber-based conductive sensing system. In other words, our approach not only enables the fabrication of individual LM fibers but also establishes a complete LM-based sensing platform. This allows for centralized sensing from single to multiple fibers, thereby broadening their applicability in electronic systems.\nIn response to the reviewer’s comments, we have revised and expanded the manuscript to better emphasize the novelty and originality of our work.\nIn the revised manuscript, “Although textile electronics have achieved considerable success in wearable applications, single fibers offer superior characteristics for particular applications6-13. ”\n“Although various solution-based methods for preparing liquid metals have been reported, the large-scale fabrication of single liquid metal fiber remains unexplored37-39. The main challenge arises from the structural differences between films and fibers: large area films possess abundant reactive interfaces that readily interact with liquid metals, whereas single fibers have a much smaller surface area and thus require higher quality electrochemical coatings to achieve uniform and stable encapsulation. Moreover, individual liquid metal electrodes are highly electrochemically active, making them unsuitable for direct application in physiological environments. To address this issue, we introduced an inert CNT/PEDOT:PSS electrode layer, allowing the exposed regions to directly contact biofluids and perform sensing functions. This bilayer electrode structure, composed of liquid metal and inert materials, combines excellent conductivity with stable sensing performance in complex biological conditions. Furthermore, by twisting and assembling multiple encapsulated liquid-metal electrodes, a multichannel sensing system can be constructed (Fig. 1b). ”\nIn the revised supplementary information,\nTable S1. Summary of recent studies on single conductive fibers\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nMelt Wearable\nspinning heater\nversatile\nintelligent\nwearable\napplications\nMelt\nspinning\nElectromagneti\nCNT/ Melt c Interference\nMXene spinning Shielding\nFabrics\nPEDOT:PSS Melt\n5 - - 3828 15 36 435\nspinning\nstretchable\nMXene/CN Melt\nT spinning\nnanogenerator\nGas sensor,\nCotto Melt\nn spinning\nsensor\nhierarchica\nElectrode,\nAgNWs/MX via lly\nene over-twisting interlocke\nsensor\nd helical\nJoule heaters,\nElectrothermoc\nPVD Coaxial\nF injected\ndisplays, and\nself-powered\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nwearable\nsensors\nOuter\nTriboelectric\nnanogenerator\nbonding\nElectrodes,\ncommunication\nfirefighting\nsuits\nStrain sensor,\nelectrode\nPDM Mold\nS injected\nOuter Physiological\nbonding Smart clothes\nStrain sensor,\nPressure\nCoaxial\ninjected\nswitches, Soft\nheater\nElectric heater,\nStrain sensing,\nCoaxial Pressure\ninjected sensing,\nProximity\nsensing\nCapillary Sustainable\ncoating\nConductive fiber\nMaximum Young’s\nRef. Conductive Raw Coating Conductivity( Diameter\nStructure Strain modulus Application\ncomponents fiber Method S/cm) (μm)\n(%) (Mpa)\nStrain sensing\napplications\nCoaxial electromagnetic\ninjected shielding\nCoaxial Physiological\ninjected Monitoring\nInductive\nEcofl Mold\nex injected\nNFC and\nMold\ninjected\nHarvesting\nStretchy Smart\nFabrics and\nCoaxial Electromagneti\ninjected c Interference\nShielding\nThi Multimodal\nFull solution Outer\ns Sensing and\nEGaIn TPU chemical layer 39623 50-100 250 1.85\nWo Stimulation\ndeposition bonding\nrk\n2. In the fabrication process of the fiber electrodes, multiple layers (e.g., PDA, AgNP, and Cu) are sequentially deposited. Each layer is expected to significantly influence the mechanical and electrical properties of the fibers. Detailed characterization of changes in fiber diameter, stretchability, and electrical conductivity at each fabrication stage should be provided.\nResponse: We sincerely thank the reviewer for this insightful comment, which we fully agree with and believe has significantly enhanced the scientific depth of our work through the supplementary analysis it prompted. In comprehensive response to the suggestion, we designed and completed a systematic series of supplementary experiments. By precisely characterizing changes in key fiber parameters before and after deposition of the polydopamine (PDA) adhesion layer, silver nanoparticle (AgNP) activation layer, copper (Cu) conductive layer, and liquid metal layer, we obtained quantitative data on the evolution of diameter distribution, mechanical properties (stress-strain curves), and electrical conductivity. The results clearly demonstrate distinct stage-specific characteristics during the transformation of the fiber from an insulating substrate to a high-performance conductive electrode: the PDA layer primarily enhances interfacial adhesion,the AgNP layer serves mainly to catalyze Cu film formation and does not provide conductivity, the Cu plating significantly improves electrical conductivity, and the liquid metal layer further enhances conductivity. Overall, the construction of each functional layer showed limited impact on mechanical properties. To visually demonstrate this multi-layer fabrication process, we have integrated a summary figure of the key data as a new chart in the revised manuscript. This figure systematically presents the stepwise diameter increase, evolution of stress-strain behavior, and orders-of-magnitude improvement in electrical conductivity, providing direct evidence for understanding the construction mechanism and performance regulation of the multi-layer electrode. We are confident that these supplementary data strongly support the rationality of the multi-layer structural design and further validate the controllability and reproducibility of the process route.\nIn the revised manuscript, “The multilayer fabrication process resulted in only a minimal increase in the overall fiber diameter, essentially adding a thin surface layer of approximately 2-3 µm (Fig. S2). Moreover, the corresponding stress-strain behavior was not significantly affected (Fig. S6). In contrast, the electrical conductivity changed markedly at each fabrication stage, as shown in Fig. S7. The initial Ag layer failed to form a stable conductive network, resulting in negligible conductivity. The subsequent Cu coating achieved a low conductivity of 0.1133 S cm⁻¹. Finally, after reaction with the liquid metal, the conductivity dramatically increased to 39623 S cm⁻¹.”\nIn the revised supplementary information,\nFigure S2. Microscopic details of the fiber at each fabrication stage. (a) Untreated TPU fiber exhibiting a smooth surface. (b) Fiber after 14 h of polydopamine (PDA) deposition, showing the formation of fine particulate structures while retaining overall smooth morphology. (c) Fiber after 3 h of AgNO₃ treatment, revealing the presence of discrete silver nanoparticles. (d) Fiber after 30 min of electrodes copper plating, displaying a dense and conglomerated copper coating. (e) Fiber after liquid metal attachment under HCl conditions, illustrating the successful binding of LM to the activated surface.\nFigure S6. Stress-strain behavior of the fiber at different fabrication stages.\nFigure S7. Electrical conductivity of the fiber at different fabrication stages.\n3. The thickness and deposition amount of the internal layers (PDA, AgNW, and Cu layer) could significantly impact the quantity and uniformity of the liquid metal coating. Is there any relevant data available?\nResponse: We sincerely appreciate the reviewer's insightful comment regarding the critical importance of controlling the thickness and quality of the inner functional layers (PDA, AgNP, and Cu) for achieving uniform and complete coating of liquid metal. In response to this valuable suggestion, we conducted a systematic optimization study of the process parameters to elucidate the influence of each functional layer on the final coating quality. Given the difficulty in directly quantifying the deposition amount of each functional layer by weight, we used reaction time as the key control parameter and performed systematic characterization of the fiber surface morphology using high-resolution SEM. This approach allowed us to establish the intrinsic relationship between process parameters, microstructure, and coating performance. Specifically, we investigated the effects of PDA polymerization time (3-14 hours), AgNP deposition time (0.5-3 hours), and Cu electroplating time (5-30 minutes) on the morphological evolution of each corresponding layer and their subsequent impact on the quality of the Galinstan coating. The experimental results demonstrate that a PDA polymerization time of 12 hours leads to the formation of a uniform and continuous adhesion layer, providing an ideal foundation for subsequent metal deposition. AgNP deposition for 3 hours achieves sufficient surface coverage, forming an effective conductive activation layer. Cu electroplating for 20 minutes produces a continuous and dense metal film, serving as an excellent conductive substrate for Galinstan coating. Together, these optimized parameters ensure ideal full-surface coverage by the final liquid metal coating.\nIn the revised manuscript, “The thickness and deposition amount of the internal layers (PDA, AgNW, and Cu) were found to be highly dependent on the deposition time. As summarized in Fig. S8, the optimized parameters are as follows: a 14-hour PDA deposition formed a uniform and continuous adhesive layer, providing an ideal foundation for subsequent metallization. A 3-hour AgNP deposition achieved sufficient surface coverage to act as an effective conductive activation layer, while a 30-minute copper electroplating produced a continuous and dense metallic film that served as an excellent conductive substrate for the subsequent Galinstan coating. Collectively, these optimized parameters ensure ideal, full-surface coverage of the final liquid metal layer”\nIn the revised supplementary information,\nFigure S8. Influence of deposition time on the morphological evolution of each functional layer. (a) A discontinuous PDA film formed after 3 h. (b) Discontinuous Ag nanoparticles appeared between 0- 2 h. (c) Corresponding elemental mapping for Ag. (d) A dense Cu layer formed within 30 min due to the fast deposition rate. (e) Corresponding elemental mapping for Cu.\n(4)The title and introduction emphasized \"scalable fabrication,\" which commonly implies suitability for mass production. However, the manuscript currently lacks experimental results whether fibers exceeding several meters can actually be fabricated, nor whether the diameter uniformity and electrical properties of the fibers can be maintained along their length. If such data cannot be provided, the use of the term \"scalable\" is not appropriate.\nResponse: We sincerely thank the reviewer for raising this important point, which gives us a valuable opportunity to further elaborate on the scalability potential of our methodology. The term \"scalable\" was carefully chosen to reflect the promising prospects of our approach for future industrial-scale applications, based on the following considerations: First, the TPU fiber production utilizes an established roll-to-roll continuous thermal stretching process (Fig. S9), which inherently possesses characteristics suitable for continuous manufacturing. Second, the liquid-phase deposition method employed in our study demonstrates excellent scalability potential in principle, as it can be readily amplified by proportionally increasing the volume of the reaction solution. Most importantly, to objectively validate this capability, we have successfully fabricated continuous TPU fibers and corresponding liquid metal fibers with lengths reaching 50 meters, with photographic evidence provided in Supporting Figure S10. While we fully acknowledge that true industrial-scale production would require addressing additional engineering challenges, these experimental results do demonstrate that our method possesses fundamental characteristics conducive to translation from laboratory-scale to industrial-scale applications.\nIn the revised manuscript, “The high throughput of the continuous melt-spinning and all-solution processes enabled the successful production of 50-meter-long TPU fibers and corresponding liquid metal-coated fibers, demonstrating their potential for scalable manufacturing (Fig. S10).”\nIn the revised supplementary information,\nFigure S10. (a) Pristine TPU fiber and (b) the resulting conductive fiber after liquid metal coating, with a total length of 50 meters.\n(5)The authors introduced additional functional layers (SEBS/CNT and PEDOT:PSS) onto the fiber electrodes. However, the electrical and electrochemical properties (Fig. 3e and Fig. 3g) are strongly dependent on the amount of PEDOT:PSS. The amount or weight% of the PEDOT:PSS should be clearly specified.\nResponse: We thank the reviewer for the valuable comment. Since the solution-based process makes it difficult to directly quantify the precise weight of PEDOT:PSS, we therefore conducted a systematic quantitative study focusing on the effect of electrochemical deposition cycles, an important parameter that directly determines the amount and thickness of the functional layer, on the electrochemical performance of the electrodes. Multiple gradients ranging from 0 to 40 deposition cycles were examined. Key parameters, including volumetric capacitance and cathodic charge storage capacity, were systematically characterized as a function of coating cycles, and cyclic voltammetry was employed to compare the electrochemical behavior of electrodes with different deposition numbers in PBS solution. The results revealed a clear saturation effect. When the number of PEDOT:PSS deposition cycles reached 30, the electrochemical capacity of the electrodes plateaued. Based on this saturation phenomenon and comprehensive performance evaluation, we ultimately selected 30 electrochemical deposition cycles of PEDOT:PSS as the optimal process condition. This choice ensures high electrochemical performance while effectively avoiding issues associated with excessive material deposition, such as increased interfacial brittleness and reduced ion diffusion rates.\nIn the revised manuscript, “Volumetric capacitance and cathodic charge storage capacity, as critical metrics for the PEDOT:PSS coating, were evaluated for electrodes with different numbers of deposition cycles (Figures S11-S12). Their electrochemical behavior in PBS solution was compared. A clear saturation effect was observed: the electrochemical capacity plateaued as the number of PEDOT:PSS deposition cycles reached 30. Based on this saturation phenomenon and a comprehensive performance assessment, 30 electrochemical deposition cycles were selected as the optimal condition.”\nIn the revised supplementary information,\nFigure S11. Electrochemical performance of the fiber as a function of the number of deposition cycles.\nFigure S12. Comparative electrochemical response in PBS solution for fibers prepared with 30 versus 40 deposition cycles.\n6. The authors demonstrated several interesting applications, such as textile-based wireless power transmission (inductor coil), ECG, and EMG signal monitoring. Specifically, the electrodes shown in Fig. 5a and Fig. 5g are designed with 2-plied fibers. Was there any signal interference (crosstalk) between the two adjacent fiber electrodes?\nResponse: We sincerely thank the reviewer for this thoughtful comment. In our design, the encapsulation layer has a thickness of approximately 25 µm, leading to a total thickness of about 50 µm when two layers are stacked. The breakdown strength of SEBS is around 39 MV/m, which means that a 50 µm layer can withstand an electric field of nearly 1950 V. Under direct current (DC) conditions, such insulation thickness is more than sufficient to prevent electrical interference. The situation is inherently more complex under high-frequency alternating current conditions. Crosstalk between adjacent conductors is primarily caused by capacitive and inductive coupling, and with an interline spacing as small as 50 µm, perceptible interference could arise in the MHz range. However, in the context of flexible sensing and electrical stimulation applications, the signals we employ are almost exclusively DC, while the AC components are of very low frequency; in this study, the maximum operating frequency is only 100 Hz. At such low frequencies, capacitive and inductive coupling effects are negligible, and significant interference is therefore unlikely. Consistently, no evidence of crosstalk was observed in our experimental studies.\n7. The authors also have presented an intriguing application involving PEDOT:PSS-coated electrodes for neural stimulation. However, PEDOT:PSS is amphiphilic and known to have potential leaching risks in wet physiological environments. Therefore, the authors should evaluate the resistance stability and amount of PEDOT:PSS under prolonged immersion conditions.\nResponse: The potential leaching of PEDOT:PSS in physiological environments is indeed a critical issue that must be addressed for its applications in bioelectronics. Compared to directly coated PEDOT:PSS films, electrochemically polymerized PEDOT:PSS coatings are generally more compact, exhibit stronger adhesion to the substrate, and contain a lower proportion of hydrophilic sulfonate groups, thereby reducing the likelihood of short-term leakage. To evaluate the stability of our PEDOT:PSS coatings, we conducted an accelerated aging test. Specifically, the fabricated fiber electrodes were continuously immersed in water at 37 °C for 72 hours, and their electrochemical responses in phosphate-buffered saline (PBS, pH = 7.4) were monitored every 12 hours. The PEDOT:PSS layer was deposited onto a robust, hydrophobic, and elastic SEBS/CNT conductive substrate, which may have facilitated strong physical and/or chemical interactions at the interface. The results demonstrated excellent stability of the PEDOT:PSS coatings. As shown in Figure S13, no significant change in interfacial impedance was observed during the 72-hour immersion period.\nIn the revised manuscript, “To ensure the reliability of the fiber electrodes in physiological environments, we evaluated the long-term stability of the PEDOT:PSS functional layer in simulated body fluid (phosphate-buffered saline, PBS, pH 7.4). Throughout the immersion period, no significant increase or fluctuation in impedance was observed, indicating that the interfacial properties between the coating and the electrolyte remained intact (Fig. S13). ”\nIn the revised supplementary information,\nFigure S13. Electrochemical response of the PEDOT:PSS electrode after immersion in PBS.\n8. When implanting electrodes in biological tissues, mechanical properties (stiffness or Young’s modulus) and weight of electrode are critically important because overly stiff or heavy electrodes can damage surrounding tissue. However, these critical mechanical parameters have not been sufficiently addressed in this manuscript.\nResponse: We thank the reviewers for their valuable suggestions. To quantitatively assess the mechanical compatibility of the fiber electrodes, we systematically measured their Young’s modulus and linear density (mass per unit length). The results indicate a Young’s modulus of 1.852 MPa and a linear density of 15.79 μg/cm, corresponding to a total mass of 1.579 mg for a 10 cm fiber. Overall, this elastic modulus is on the same order of magnitude as that of neural tissue, slightly higher than the modulus of the nerve. It is noteworthy that similar polymer materials, including the same type of TPU used in this study, have been extensively validated for in vivo implantation and neural stimulation applications (Advanced Functional Materials, 2023, 33(33): 2303134; Neural Networks, 2023, 162: 212-224; Advanced Sensor Research, 2025, 4(7): 2500002; Advanced Healthcare Materials, 2025, 14(3): 2404296), demonstrating that materials with comparable mechanical properties and weight can be safely employed in implantable electronic devices.\nIn the revised supplementary information,\nFigure S14. Young's modulus and mass density of the conductive fiber.\n9. Overall, the manuscript is well-written. However, minor typographical errors remain (e.g., insperion, form Covestro, etc).\nResponse: We thank the reviewer for the nice reminder. We double checked the manuscript. The errors have been fixed in the revised manuscript."}, {"round": 2, "reviewer_comments": {"reviewer_1": "The authors have addressed all my concerns and I recommend publishing this work.", "reviewer_2": "I really appreciate the author's great efforts in addressing my concerns. Now, I have no question about the experiment details in their work.\nHowever, I still have critical issues about this manuscript.\n1. Although the author stated that they provided a study about a soft conductor, the applications used the fibers to collect ECG signals on skin surface (Figure 5) and recognized muscle contractions in vivo (Figure 6). The authors actually utilized the fibers as sensors. That's why I stressed that the article worked on a soft sensor since the last two applications were very typical sensoring set-up experiments. So, I can not agree with the perspective in the response letter that they offered a soft conductor.\n2. Besides, if they provide a soft conductor, the performance of the conductor is poor. The R/R0 is very large (reaching 4 at 40% strain), which can not persuade me that this was an excellent soft conductor.\n3. I think the manuscript is not clear about sensor or conductor. Sensors are used to detection but conductor should provide the stable electrical conductivity under extremely large strain (e.g. R/R0 < 1.5 under 600% strain).\n4. The author should re-write their manuscript to clarify their major contribution in electronics. In current form, the topic and final demos are very confusing.", "reviewer_3": "The authors respond to the reviewers' questions thoroughly, with sufficient detail and appropriateness."}, "author_response": "Point-by-point responses to the reviewers’ comments\nDear Reviewers,\nThanks for your consideration for publication of our work entitled \"Scalable and Stretchable 1D Bioelectronic Fibers for Multimodal Sensing and Stimulation\" (Manuscript ID: NCOMMS-25-48993A-Z).\nWe are very grateful for the detailed and thoughtful comments made by the reviewers. The comments are very helpful for us to further improve our work. The revision has been made to fully address the reviewer’s comments point by point. All revisions have been highlighted in red color in the revised manuscript for the convenience of reviewing.\nWe thank you in advance for your time and attention.\nWith Warm Regards,\nYours sincerely,\nYunlei Zhou\nReviewer #1\nThe authors have addressed all my concerns and I recommend publishing this work.\nResponse: We sincerely appreciate the reviewer's positive evaluation of our work. We are delighted that all concerns have been thoroughly addressed and are grateful for the recommendation for publication. Thank you once again for your valuable insights and support throughout the review process.\nReviewer #2\nI really appreciate the author's great efforts in addressing my concerns. Now, I have no question about the experiment details in their work.\nHowever, I still have critical issues about this manuscript.\nResponse: We sincerely thank you for your continued guidance and insightful comments. We are glad that the previous revisions addressed your concerns on the experimental details, and we have given our utmost attention to your remaining critical points on other aspects of the manuscript. To address these, we have further revised the manuscript to enhance its clarity, rigor, and overall presentation. Our point-by-point responses and the corresponding revisions are detailed below. We trust these revisions will meet your approval and enhance the quality of our work.\n1. Although the author stated that they provided a study about a soft conductor, the applications used the fibers to collect ECG signals on skin surface (Figure 5) and recognized muscle contractions in vivo (Figure 6). The authors actually utilized the fibers as sensors. That's why I stressed that the article worked on a soft sensor since the last two applications were very typical sensoring set-up experiments. So, I can not agree with the perspective in the response letter that they offered a soft conductor.\nResponse: We sincerely appreciate your insightful comment and fully agree with your observation. Our earlier description of the device as a soft conductor was indeed insufficiently accurate, and we apologize for the lack of clarity. In fact, the device is fabricated by first creating a liquid metal based conductive fiber, followed by forming an electrochemically stable PEDOT:PSS layer at the designated sensing region to ensure stable sensing performance. In our previous response, our intention was simply to emphasize that the sensing behavior originates from the material’s intrinsic conductive properties rather than from strain-dependent mechanisms. However, we recognize that our wording may have caused misunderstanding, and we are grateful for the opportunity to clarify this point.\n2. Besides, if they provide a soft conductor, the performance of the conductor is poor. The R/R0 is very large (reaching 4 at 40% strain), which cannot persuade me that this was an excellent soft conductor.\nResponse: We fully agree with your point that the bilayer electrode should not be regarded as a simple conductor. Our intention was only to clarify that the device does not operate as a resistive type strain sensor but rather as an electrode for signal transduction. We also appreciate your concerns regarding the electrode performance. In fact, PEDOT is a rigid conjugated polymer chain with very limited intrinsic stretchability. The strong coulomb interactions and the hydrogen bond network between PEDOT and PSS make the material brittle, and typical unmodified PEDOT:PSS exhibits conductive pathway failure at only 2–10% strain, resulting in a sharp increase in resistance. Through our composite electrode design, we enhance the stretchability of PEDOT:PSS from 2–10% to over 100%. Although many studies have explored PEDOT:PSS modification, such as forming co-continuous structures with SEBS, PU, or hydrogels to maintain conductivity at strains exceeding 100%, these approaches often compromise the intrinsic conductivity, which is typically around 10 S cm⁻¹ (e.g., Nat. Mater. 22, 895–902, 2023). Furthermore, the electrochemical impedance of the stretched electrode shows minimal deviation from its original state, despite a substantial increase in measured resistance. This is because the electrode primarily serves as a passive interfacial contact rather than an active component in the electrochemical reaction. Overall, our work provides a bilayer electrode platform that integrates both stretchable conductivity and electrochemical stability, enabling stable sensing, electrical stimulation, and other bioelectronic applications, while also offering the advantage of a small overall fiber size.\n3. I think the manuscript is not clear about sensor or conductor. Sensors are used to detection but conductor should provide the stable electrical conductivity under extremely large strain (e.g. R/R0 < 1.5 under 600% strain).\nResponse: We sincerely appreciate your insightful comment. Our intention was to clarify that we first fabricated a single liquid metal based conductive fiber, upon which an inert sensing electrode was subsequently constructed. Although liquid metal conductors are known to maintain stable electrical conductivity under extremely large strains (for example, R/R₀ < 1.5 under 600% strain), achieving such stability in a single liquid metal fiber remains highly challenging. For instance, a recently reported fiber composed of thermoplastic polyurethane, micron sized silver flakes, and liquid metal (Nature Communications, 2025, 16, 6785) exhibited an R/R₀ greater than 200 under 300% strain. Certain liquid metal fibers exhibit improved electrical stability, such as those fabricated with a core shell structure in which the liquid metal is fully encapsulated. However, in these cases the liquid metal cannot participate in sensing because it is entirely isolated within the elastomer matrix. In addition, our fiber is extremely thin, and at such a small diameter even liquid metal can exhibit a noticeable increase in R/R₀ under strain (Nature Communications, 2025, 16, 7920). Overall, the conductivity of our fiber electrode remains within an acceptable and functional range. Moreover, by integrating the liquid metal conductor with the composite sensing electrode, we demonstrate that the resulting device not only serves as a conductor but also enables stable sensing performance in electrochemical environments. Although the resistance under tensile strain increases markedly, this effect is primarily associated with the contact interface; as a result, only negligible changes in electrochemical impedance are observed between the stretched and pristine states in practical electrochemical measurements.\n4. The author should re-write their manuscript to clarify their major contribution in electronics. In current form, the topic and final demos are very confusing.\nResponse: Thank you very much for your valuable comment. After a systematic reassessment of the research content and careful consideration of your suggestions, we have restructured the manuscript to enhance its logical coherence and academic rigor. The central objective of this work is clearly defined as the development of a scalable and stretchable one-dimensional multifunctional fiber capable of multimodal electrical signal transmission, sensing, and electrical stimulation. To highlight this main theme, the overall organization of the manuscript has been revised into a logical framework that progresses from device design to fabrication and finally to validation across multiple functional scenarios.\nThe manuscript title has been updated to Scalable and Stretchable 1D Multifunctional Fibers for Multimodal Sensing and Stimulation. To enable readers to understand the structural configuration and design rationale of the fiber at the outset, the original Figure 1, which illustrates the overall design of the stretchable multifunctional one-dimensional fiber, has been moved to the Introduction section. Within the Results section, the previous subsections “Fabrication and Characterization of Stretchable Liquid Metal Fiber” and “Seamlessly Integrated with Stretchable Electrodes” have been combined into a unified subsection titled “Fabrication and Characterization of Stretchable Multifunctional Fiber.” This revised subsection presents a comprehensive description of the fabrication workflow, material assembly, and key performance characterizations of the fiber, thus providing a consistent technical foundation for the subsequent application-focused sections.\nThe three application-related subsections have been reorganized to emphasize the multifunctionality of the fiber across different operational contexts:\nIntegration of Stretchable Multifunctional Fiber into Textiles for Wireless Energy Transmission. This section demonstrates the fiber’s capability as a stretchable conductive component for wireless power transfer within flexible textile systems.\nApplication of Stretchable Multifunctional Fiber in Electrophysiological Monitoring. This section validates the fiber’s reliability and effectiveness as wearable sensors for ECG and EMG signal acquisition.\nApplication of Stretchable Multifunctional Fiber for In Vivo Nerve Stimulation. This section evaluates the fiber’s functionality and stability as an implantable electrical stimulation device within electrochemically active in vivo environments.\nThese revisions, carried out to the best of our abilities, are intended to further enhance the clarity, coherence, and overall scholarly quality of the manuscript. By establishing a continuous narrative that progresses from device design to fabrication and characterization, and subsequently to representative application demonstrations, the revised structure seeks to more clearly convey the practical significance and broad potential of the proposed stretchable multifunctional fiber in wireless energy transmission, wearable electrophysiological monitoring, and in vivo neural stimulation. In addition, we have further refined the Introduction, Conclusion, and other specific sections of the manuscript to improve clarity and accuracy. Given the comprehensive nature of these revisions, the detailed changes are not enumerated here but can be directly reviewed in the updated manuscript. We sincerely hope that these revisions meet your expectations, and we are deeply grateful for your insightful, patient, and generous guidance, which has been invaluable in enhancing the overall quality and rigor of this work.\nReviewer #3\nThe authors respond to the reviewers' questions thoroughly, with sufficient detail and appropriateness.\nResponse: We sincerely appreciate the reviewer’s positive comments. We are pleased to learn that our responses were considered thorough, sufficiently detailed, and appropriate. Thank you for your careful evaluation and constructive feedback throughout the review process."}]

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