Highly transparent, intrinsically stretchable, photo-patternable, and vacuum-deposited electrodes for wearable sensors and displays

Abstract

The development of stretchable and transparent electrodes is essential for next-generation wearable displays, human–machine interfaces, and on-skin bioelectronic devices; however, conventional approaches are limited by low fabrication compatibility with conventional semiconducting manufacturing processes, unstable electrical conductivity under stretching, and limited non-uniform areal transparency. Here, we report a novel device fabrication strategy for developing a highly transparent, intrinsically stretchable, photo-patternable, and vacuum-deposited (T-iSPV) electrode. The strain-insensitive performance of the T-iSPV is inherent in the in situ formation of a conducting bilayer consisting of a crack-based Au nanomembrane and Au–elastomer nanocomposite during direct thermal deposition of Au onto an elastic substrate. In addition, a photo-patterning process and optimal thickness/design and evaporation rate of the Au bilayer delicately balance the stretchability, electrical conductivity, and transparency of the T-iSPV. To demonstrate its versatility, the T-iSPV is applied as a conformal bioelectronic interfacing electrode for monitoring electrocardiogram (ECG), electromyogram (EMG), and electrooculogram (EOG) signals. Furthermore, the T-iSPV electrochemically activates the stretchable active layers composed of poly-3-hexylthiophene (P3HT) in a styrene–ethylene–butylene–styrene (SEBS) polymer matrix to effectively modulate electrochromic displays. These findings underscore the potential of the T-iSPV for enabling the evolution of next-generation conformal bioelectronic and optoelectronic systems.

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New concepts

This study introduces novel highly transparent, intrinsically stretchable, photopatternable, and vacuum-deposited (T-iSPV) electrodes that simultaneously achieve stretchability, transparency, and stable electrical performance, offering a breakthrough in bioelectronics and wearable optoelectronic systems. Unlike conventional transparent stretchable electrodes, which are often limited by fabrication inconsistencies (i.e., spray-coated metal nanowires) or poor electrical conductivity (i.e., organic conductive materials), our approach leverages a precision-engineered Au mesh pattern on an ultrathin styrene–ethylene–butylene–styrene (SEBS) substrate fabricated via a photopatterning lift-off process. This method effectively eliminates undesired nanoparticle infiltration into SEBS, improves optical transparency while maintaining mechanical robustness, and prevents potential electrical crosstalk. The distinguishing feature of this approach is its ability to integrate passive and active bioelectronic functionalities. By fine-tuning the design parameters of the mesh structure, we systematically modulated the stretchability and conductivity to optimize its performance for diverse applications, including transparent electrocardiogram and electromyogram electrodes and optoelectronic components such as electrochromic displays. This discovery not only addresses key challenges in wearable and implantable devices, but also sets a precedent for the development of next-generation electronic skin with real-time biosignal monitoring and display functionalities.

Introduction

Wearable biointegrated electronics employing flexible and stretchable functional devices have emerged as a promising technology for monitoring physiological signals,^1–4 transmitting real-time health data,^5–8 and enabling precise diagnostics through the adoption of machine learning techniques based on artificial intelligence^9–11 and personalized therapy in a closed-loop manner for disease management.^12–14 To perform these diverse functions, various electronic components must be incorporated into wearable devices, including highly sensitive biosensors for precise signal acquisition,^15–18 displays for visualizing information,^19,20 and analog front-end circuits for processing microelectronic signals^21,22 and energy storage devices.^23,24 These functional elements should also possess soft mechanical properties and conform to the curvilinear and dynamic nature of the human skin.^25–28 However, conventional electronic circuit components based on rigid inorganic materials struggle to achieve conformal contact with the skin and lack mechanical deformability, leading to discomfort and limited usability.²⁹

To overcome these challenges, soft, stretchable conductive materials have been explored for long-term use in stable wearable electronics. Ultrathin stretchable electronics have garnered significant attention owing to their unique ability to conform closely to various body locations while remaining mechanically imperceptible to the wearer.^30,31 In addition to mechanical compliance, visual imperceptibility is crucial for enhancing the wearability of such devices from an aesthetic point of view, necessitating the development of stretchable and transparent electrodes.^32,33

Indium tin oxide (ITO) is widely used as a transparent electrode; however, its inherent brittleness makes it unsuitable for applications requiring mechanical deformation. To address this issue, alternative stretchable transparent electrodes have been developed using organic conductive polymers, such as poly(2,3-dihydrothieno-1,4-dioxin)-poly(styrenesulfonate) (PEDOT:PSS),^34,35 and carbon-based conductive materials, such as single-walled carbon nanotubes (CNTs)^36,37 and graphene,^38,39 incorporated into stretchable polymer substrates. Despite their transparency, these conducting materials exhibit poorer conducting properties than metallic materials, limiting their efficiency.⁴⁰ Metal nanowires, such as silver nanowires (AgNWs), have also been explored to achieve high electrical conductivity in stretchable electrodes;^41–43 however, fabrication techniques such as spray coating introduce challenges related to process uniformity, high manufacturing costs, and incompatibility with semiconductor processing.^44,45 To overcome these limitations, metal mesh structures have been proposed, resulting in high compatibility with conventional fabrication processes and high areal uniformity.^46,47 However, a critical challenge remains: wavy pattern designs tailored to improve stretchability significantly degrade transparency. In this regard, the development of an intrinsically stretchable electrode with high transparency is essential; however, such an electrode has not yet been developed.

Herein, we report a new class of flexible conductors, a highly transparent, intrinsically stretchable, photopatternable, and vacuum-deposited (T-iSPV) electrode, fabricated by thermally evaporating Au nanoparticles (AuNPs) onto an ultrathin styrene–ethylene–butylene–styrene (SEBS) substrate with a modified mesh-patterned structure (Fig. 1a, left). According to previous reports, Au nanoparticles (AuNPs) can penetrate the SEBS matrix during the thermal evaporation process.⁴⁸ As a result, when the mesh pattern was fabricated through a subsequent wet etching process, the AuNPs embedded within the SEBS were not completely removed, leading to reduced transparency (Fig. 1a, center and top). To address this issue, a photoresist (PR) was first photopatterned onto the SEBS substrate, and the Au nanomembrane was then thermally deposited onto the SEBS. The photopatterning lift-off process was applied to eliminate undesired regions of Au deposition (Fig. 1a, center and bottom). This approach enabled the formation of a mesh pattern without residual AuNPs, thereby enhancing transparency (Fig. 1a, right).

Fig. 1 Overview of the transparent, intrinsically stretchable, photopatternable, and vacuum-deposited (T-iSPV) electrode. (a) Schematic illustration of the T-iSPV electrode (left), the fabrication process of the control electrode (center, top) and the T-iSPV electrode (center, bottom), and a photographic image of the T-iSPV electrode (right). (b) Schematic depiction of AuNPs penetrating the SEBS layer depending on the thermal evaporation rate and the expected microcrack size of the Au nanomembranes. (c) Schematic illustration showing changes in the stretchability and transparency of the T-iSPV electrode under different widths and pitches. (d) Optical microscope (OM) images of the T-iSPV electrode on SEBS substrates of different thicknesses: 200 μm (top) and 5 μm (bottom) attached to the skin. (e) Schematic illustration of the integrated wearable bioelectronic system comprising a tactile sensor, electrochromic display, and electrocardiogram electrode based on T-iSPV. (f) Photograph of the integrated wearable bioelectronic display system attached to the skin (left) and under compressive and tensile strain (right). [Illustration in (c) was created with BioRender.com].

The design rules for improving the transparency and stretchability of the T-iSPV with respect to the Au deposition conditions and pattern designs are given in Fig. 1b and c. As shown in Fig. 1b, the depth of AuNP penetration into SEBS depends on the thermal evaporation rate, which in turn affects the size of the microcracks formed upon deformation. As expected, the size of these microcracks on the SEBS determined the resolution of the micropattern and the overall stretchability of the T-iSPV electrode. Additionally, as illustrated in Fig. 1c, the transparency and stretchability of the T-iSPV were influenced by the pitch and linewidth of the mesh pattern. Transparency is dictated by the ratio of the pitch to the line width, whereas stretchability increases with a wider line width. However, when the pitch exceeds 100 μm, it surpasses the visual perception threshold, resulting in decreased transparency.⁴⁹ Following these design rules, the T-iSPV electrode was successfully used for conformal attachment to the skin, verifying its potential application in wearable transparent electronics (Fig. 1d–f).

In this context, the T-iSPV electrodes are highly suitable for high-performance wearable bioelectronic/optoelectronic applications owing to their superior stretchability and transparency (Tables S1 and S2, ESI†).

Experimental

Materials

Octadecyltrichlorosilane (OTS), chloroform, and acetone were purchased from Sigma-Aldrich. Regioregular poly(3-hexylthiophene-2,5-diyl) (P3HT) (Mn: 54 000–75 000, lot number: MKCK1947), poly(vinylidene fluoride-co-hexafluoropropylene) (PVDF-HFP) (Mn: 110 000), and the ionic liquid 1-ethyl-3-methylimidazolium bis(trifluoromethyl sulfonyl)imide (EMIM:TFSI) were also purchased from Sigma-Aldrich. SEBS copolymers H1062 and H1221 were purchased from Asahi Kasei, and polydimethylsiloxane (PDMS) was acquired from Dow.

Fabrication of the conducting bilayer consisting of a crack-based Au nanomembrane and Au–elastomer nanocomposite

A SEBS elastomer of 180 mg mL⁻¹ in toluene (H1062, Asahi Kasei) was spin coated onto an octadecyl trimethoxysilane (OTS)-treated wafer (or glass) (1000 rpm/1 min, resulting in 5 μm in thickness). Gold (thickness of 50 nm) was thermally evaporated on the SEBS substrate at a deposition rate of 5.0/0.5/0.1 Å s⁻¹.

Microfabrication of T-iSPV using the lift-off process

A positive photoresist (S1805) was patterned on the SEBS substrate using a negative photomask. Then, Au (thickness of 50 nm) was thermally evaporated on the photoresist-patterned substrate at a deposition rate of 0.1 Å s⁻¹. The AuNPs selectively diffused into the uncovered SEBS surface. Finally, the Au layer on the photoresist was lifted-off to obtain micro-cracked Au nanomembranes and Au–elastomer nanocomposites.

Electrochemical impedance spectroscopy measurement

Electrochemical impedance spectroscopy (EIS) was performed on the T-iSPV electrode using a potentiostat (ZIVE sp1, ZIVE LAB). The electrode, with an active area of 1.0 × 1.0 cm², was connected as the working electrode. An Ag/AgCl electrode served as the reference, and a platinum electrode was used as the counter electrode. Impedance measurements were performed over a frequency range of 0.1 Hz to 10 kHz.

Electrophysiological signal measurement

Before conducting the experiment, the complete procedure was approved by the Institutional Review Board (IRB) of Sungkyunkwan University (Approval No. SKKU 2025-02-082).

Electrophysiological signals, including ECG, electromyography (EMG), and electrooculography (EOG), were measured using T-iSPV electrodes attached to human skin with a Tegaderm film (3M, Saint Paul, MN, USA). For a comparative analysis, conventional Ag/AgCl electrodes were simultaneously employed. Signals were acquired using a bio-signal amplifier (Bio Amp, AD Instruments, Dunedin, New Zealand) and a data acquisition device (PowerLab 8/35, AD Instruments, Bella Vista, Australia) at a sampling frequency of 1 kHz, and processed using LabChart 8 Pro software (AD Instruments, Bella Vista, Australia).

ECG signals were recorded using a standard three-lead setup, with the reference electrode placed on the right wrist, active electrode on the left wrist, and ground electrode on the left ankle. ECG measurements were performed for ≈10 min in a resting state. Additionally, ECG signals were measured before and after performing 50 jumping jacks, each session lasting ≈5 min, to evaluate performance during physical activity. ECG signals were filtered using a 0.5–200 Hz bandpass filter to eliminate unwanted noise.

EMG signals were recorded using T-iSPV electrodes fixed to the skin of the left forearm using a Tegaderm film. A conventional electrode placed on the ankle served as the ground electrode, and both electrodes were positioned at regular intervals along the forearm. EMG measurements involved two distinct movements: grasping an object and lifting the wrist. EMG signals were bandpass filtered from 30 to 1000 Hz according to standard EMG measurement guidelines.

For the EOG measurements, conventional electrodes were placed above and below the left eye, and the T-iSPV electrodes were similarly placed around the right eye. A conventional electrode positioned between the eyebrows served as the ground. EOG signals were processed with a 0.1–30 Hz bandpass filter.

All recorded electrophysiological signals underwent additional processing using a 59–61 Hz bandstop filter to eliminate line-noise interference. Subsequently, the signals are denoised using a 1D wavelet-denoising algorithm. The maximum overlap discrete wavelet transform method was applied to ensure the robustness of wavelet-based denoising.

Preparation and characterization of a stretchable electrochromic display (sECD)

To prepare the substrate, a Si wafer was treated with O₂ plasma (100 W, 5 min) and subsequently immersed in an OTS solution (0.5 wt% in n-hexane). After 1 h, the substrate was rinsed with ethyl alcohol and heated at 120 °C for 30 min. To prepare the P3HT solution, 13 mg of P3HT and 7 mg of SEBS-H1221 were dissolved in 1.2 mL of chloroform. The P3HT solution was then spin-coated onto the OTS-treated substrate at 1000 rpm for 1 min. The ion gel solution was prepared by blending PVDF-HFP, EMIM:TFSI, and acetone in a weight ratio of 1 : 4 : 7, followed by stirring at 60 °C for 6 h. To fabricate the sECD, the P3HT/SEBS composite film was transferred onto the bottom T-iSPV electrode and annealed at 60 °C for 5 min under applied pressure to ensure proper adhesion between each layer. The ion gel solution was spin-coated onto the top T-iSPV electrode at 1000 rpm for 1 min. Then, the solid electrolyte on the contact pad of the bottom electrode was removed using a razor blade. To integrate the sECD, the ion gel-coated bottom T-iSPV electrode and the P3HT/SEBS-transferred top T-iSPV electrode were brought into contact and annealed at 60 °C for 5 min under applied pressure.

Transparent integrated bioelectronic display system demonstration

To reduce contact resistance, liquid metal was applied between the electrodes and the contact pads of the sECD and tactile sensor, allowing the integration of copper wires. For stable attachment on the skin, the sECD and ECG electrodes were first affixed onto medical tape (Tegaderm, 3M), which was then applied to the arm. The tactile sensor was assembled on the top surface of the attached medical tape and secured with a second layer of medical tape. The tactile sensor was connected in parallel with a 3 MΩ reference resistor to the analog ground terminal of an Arduino Mega 2560. For measurement, an additional 3 MΩ resistor was connected in series between the sensor and a 5 V input, and the voltage across the sensor was measured through the analog input of the Arduino. The sECD was connected to a pulse width modulation (PWM) pin of the Arduino, and a ±2 V alternating polarity voltage was applied whenever the toggle signal was activated. The ECG electrodes were connected to a data acquisition system (PowerLab 8/35, AD Instruments, Bella Vista, Australia) with a sampling frequency of 1 kHz and processed using LabChart 8 Pro software (AD Instruments, Bella Vista, Australia). Both the Arduino and DAQ system were controlled using a customized MATLAB code.

Results and discussion

Fabrication and characterization of T-iSPV

To understand the nature of the conducting bilayer consisting of a crack-based Au nanomembrane and Au–elastomer nanocomposite, we investigated its surface and inner structures at the nanoscale level. Optical microscopy (OM) and scanning electron microscopy (SEM) mapping revealed the formation of microcracks on the Au nanomembrane surface (Fig. 2a, right). The microcracks were clearly distinguishable under mechanical strain. Microscopic mapping revealed that the AuNP nanodispersion affected the microscopic electrical and mechanical properties of the Au nanomembrane. The formation of the Au–elastomer nanocomposite resulted in high stretchability, even after micropatterning, whereas the lack of it made the Au nanomembrane non-stretchable (Fig. 2a, left and center). On a microcracked Au nanomembrane, the ratio of nano-dispersed AuNPs to elastomer is optimal at a deposition rate of 0.1 Å s⁻¹; the Au–elastomer nanocomposite adheres the Au nanomembrane to the elastomer surface and forms microcracks. As an Au–elastomer nanocomposite containing a deep-percolated network of elastomers and AuNPs under the surface of the Au nanomembrane, microcracks formed on the surface and the conductive connection remained stable under mechanical peeling and strain, even after stretching cycles (Fig. 2b, c and Fig. S1–S3, ESI†).

Fig. 2 Fabrication and characterization of T-iSPV. (a) OM and SEM mapping show the surfaces of the non-stretchable electrode (left, 5.0 Å s⁻¹), non-micro-patternable electrode (center, 0.5 Å s⁻¹), and microcracked Au nanomembrane electrode (right, 5.0 Å s⁻¹). (b) Resistance changes in Au nanomembrane electrodes under strain. The sample has a width of 1 cm and a length of 3 mm. (c) Resistance changes in Au nanomembrane electrodes during 30% stretching cycles. The inset shows magnified resistance changes over 20 stretching cycles. (d) OM image of the as-fabricated T-iSPV electrode after the lift-off process. (e) Sheet resistance and transmittance of T-iSPV as determined by the ratio of pitch to line width. (f) and (g) Schematic illustrations of the fabrication processes for wet etching (f) and lift-off (g). (h) Photographs of the wet-etched control electrode (left) and the T-iSPV electrode (right). (i) Schematic showing the blank region of the control (left) and T-iSPV (right). The corresponding atomic concentration from TOF-SIMS indicates that AuNPs are diffused within the SEBS elastomer matrix. Au (yellow) and C (blue) were selected as the characteristic groups during sputtering to evaluate the nano-dispersion of the AuNPs into the SEBS elastomer matrix along the film thickness direction. (j) Photographs of the freestanding ultrathin T-iSPV at different tensile strains. (k) Resistance changes in T-iSPV electrodes under strain. The inset shows a microcracked T-iSPV. (l) Normalized transmittance changes in T-iSPV electrodes under strain.

Thermally deposited Au on an elastomer substrate with a thickness of 40 nm usually has a low transmittance, but our photopatterned Au meshes are highly transparent (Fig. S4, ESI†). The remarkable performance of the Au mesh electrodes can be attributed to several factors. First, Au meshes were fabricated using a standard vacuum-deposition process before the photopatterning lift-off process, which yielded sharply defined, highly uniform, and reproducible electrode geometries (Fig. 2d and Fig. S5–S7, ESI†), which is comparable to those of spray-coated metal nanowires (i.e., m-CNTs, AgNWs, and PEDOT:PSS). Notably, nanostructured materials usually exhibit significantly unstable conductivities owing to fabrication inconsistencies. Second, the Au meshes form a highly uniform, interconnected network determined by the pitch (p) and line width (w), according to the following equation:⁵⁰

T = (p − w)²/p²

The T-iSPV mesh electrodes must possess a theoretical transparency of 76.6% by maintaining a pitch and width ratio of 8, while ensuring consistent electrical resistance (Fig. 2e and Fig. S8, ESI†). These Au meshes must have exceptional parameters in order to be applied in high-performance transparent conducting electrodes. The resistance has been shown to largely depend on the line width of the network owing to the greater influence of microcracks in the Au nanomembrane, whereas the transmittance remained at around 80%, regardless of the pitch and width dimensions. Therefore, the improvement observed here can be ascribed to the visual imperceptibility resulting from the human visual perception threshold achieved by not exceeding a 100 μm pitch (Fig. 2e, inset). The precisely interconnected structure of the Au meshes consisting of crack-based Au nanomembranes and Au–elastomer nanocomposites (Fig. 1a, right, and Fig. 1d) also prevents the creation of a large junction resistance, which is a common bottleneck in metal–nanowire networks.⁵¹

The fabrication of the T-iSPV was based on a conventional photopatterning process. An Au nanomembrane on an elastomer substrate requires precise optimization owing to its deformable nature near the glass transition temperature (T_g).^52,53 In this study, we fabricated microcracked Au meshes that are photolithography compatible and do not require chemical or structural engineering. As shown in Fig. 2f, in the case of Au deposited on the SEBS substrate followed by PR patterning and wet etching, the nano-dispersed AuNPs remained embedded in the SEBS, reducing the transparency and necessitating their removal. In contrast to the AuNPs penetrating the surface of the bulk elastomer matrix and diffusing throughout the free volume after the wet etching process, the vacuum deposition of Au on the selectively opened surface of SEBS resulted in more uniform transparency, achieving a transmittance of 77.17% at a 550 nm wavelength (Fig. 2g, h and Fig. S9, ESI†). The AuNPs nano-dispersed between the lines were characterized using a secondary ion mass spectrometer (SIMS). The SIMS depth profile confirmed that the residual diffusion of AuNPs in the undesired region of the mesh after the wet etching process showed gradual diffusion of the Au atomic concentration from the surface of SEBS, whereas the T-iSPV consistently demonstrated no atomic concentration of Au, indicating the effective transmittance of the Au mesh (Fig. 2i and Fig. S10, ESI†).

The T-iSPV electrode was both conformable and robust under strain conditions. It can be manipulated effectively during fabrication or at the skin-surface interface (Fig. 1d, bottom, and Fig. S11, ESI†). Owing to the intrinsically stretchable properties of the bilayer consisting of crack-based Au nanomembranes and Au elastomer nanocomposites, the T-iSPV electrode exhibited both high stretchability and transparency. Even on an ultrathin substrate (5 μm), the T-iSPV electrode maintained its transparency while withstanding over 50% strain (Fig. 2j–l and Fig. S12, S13, ESI†). In addition, the T-iSPV electrode exhibited stable conductivity after cycle stretching within the typical range of skin deformation (Fig. S14, ESI†). Together, these experiments demonstrate that the conformal contact and durable electrical and mechanical properties of the T-iSPV electrodes achieved through the crack-based Au nanomembranes and Auelastomer nanocomposite structures enable delicate electrophysiological signal recording and subsequent display applications. This is not possible using conventional stiff bulk networks with low transparency, which hardly achieve electrical connectivity or conform well to the skin surface.

Electrophysiological monitoring using the T-iSPV electrode

A T-iSPV electrode was used to measure the representative electrophysiological signals from the skin, including ECG, EMG, and EOG. For a direct comparison, simultaneous measurements were performed using conventional Ag/AgCl electrodes. The overall monitoring setup, employing both the T-iSPV and conventional electrodes, is illustrated in Fig. 3a.

Fig. 3 Electrophysiological monitoring using the T-iSPV electrode. (a) Overall scheme for electrophysiological signal monitoring using the T-iSPV and control electrodes. (b) Photograph showing the attachment of T-iSPV and control electrodes on the forearm for ECG and EMG measurement. (c) and (d) ECG signals monitored before and after physical exercise using the T-iSPV (blue) and control electrodes (black). (e) PQRST complex of ECG signals recorded using T-iSPV (top) and control electrodes (bottom). (f) Corresponding signal-to-noise ratio (SNR) analysis of ECG data (n = 3). Unpaired two-tailed t-test: NS, not significant, P = 0.15. (g) Photograph showing wrist movement during EMG measurement. (h) and (i) EMG signals recorded during wrist movements using T-iSPV (blue) and control electrodes (black). (j) Corresponding SNR comparison of EMG data during wrist movements. Unpaired two-tailed t-test results indicated that during grab movements, there was NS (P = 0.071), whereas raise movements showed a statistically significant difference (*P = 0.015). (k) Photograph showing the attachment of T-iSPV and control electrodes on the forehead for EOG measurement. (l) EOG signals recorded during eye movements with T-isPV and control electrodes. (m) Power spectrogram of EOG signals corresponding to individual eye movements recorded using the T-iSPV electrode. [Illustration in (a) and (l) were created with BioRender.com].

For ECG recordings, both electrode types were attached to the wrists according to the three-lead pacing method, with an additional conventional electrode on the ankle serving as the ground. In the EMG measurements, the conventional electrode on the ankle functioned as the ground, whereas the T-iSPV and conventional electrodes were placed at regular intervals along the left wrist. For the EOG recordings, conventional electrodes were placed above the left eye, and the T-iSPV electrodes were similarly positioned around the right eye. Another conventional electrode, located on the left side of the neck, acted as the ground. The T-iSPV electrode was secured to the skin using 3M Tegaderm.

Signals were acquired at a sampling rate of 1 kHz using a data acquisition system (PowerLab 8/35; AD Instruments) combined with a BioAmp amplifier (AD Instruments). ECG and EMG signals were denoised and filtered to eliminate line noise and environmental disturbances, while EOG signals were bandpass filtered within the 0.1–30 Hz range.

The impedance of the T-iSPV electrode was measured across frequencies from 0.1 Hz to 10 kHz, corresponding to the frequency bands of ECG, EMG, and EOG signals. The results reveal that as the frequency decreases, the impedance increases—from 843 Ω at 10 kHz to 22 349 Ω at 100 Hz (Fig. S15, ESI†).

Fig. 3b shows conventional and T-iSPV electrodes attached to the arm for ECG and EMG monitoring. The ultrathin and transparent characteristics of the T-iSPV electrode allowed it to closely conform to the skin and preserve skin visibility (Fig. S16, ESI†). ECG measurements were initially conducted for 5 min at rest, followed by a session of 50 jumping jacks, after which measurements were recorded immediately. Before the exercise, both electrode types demonstrated comparable signal-to-noise ratios (SNRs) (Fig. 3c), with a resting heart rate of ≈70 BPM. Immediately after exercise, the heart rate increased to nearly 100 BPM. Moreover, the PQRST waveforms were clearly discernible with both electrodes (Fig. 3e), and the SNR values were statistically similar (ns, P = 0.15), indicating comparable ECG measurement performance for the T-iSPV and conventional electrodes (Fig. 3f).

For EMG monitoring, two conventional and two T-iSPV electrodes were placed at regular intervals along the forearm while subjects performed object-grabbing and raising movements (Fig. 3g). Both electrode types exhibited comparable EMG signal amplitudes during object-grasping and wrist-lifting actions (Fig. 3h and i). During the object-grabbing task, both electrode types yielded similar SNR values. During wrist raising, however, the T-iSPV electrode demonstrated a slightly superior SNR, indicating more reliable and stable EMG signal detection (Fig. 3j).

EOG measurements involved recording the right eye movements with conventional electrodes and the left eye movements with T-iSPV electrodes (Fig. 3k). Both electrode types effectively detected eye-movement signals, exhibiting opposite polarities owing to the different grounding positions (Fig. 3l). Similar amplitude variations confirmed the comparable EOG measurement capabilities of the conventional and T-iSPV electrodes. Moreover, similar amplitude values were observed for each movement, confirming the comparable EOG measurement capabilities of the conventional and T-iSPV electrodes. Power spectrogram analysis of the EOG signals acquired with the T-iSPV electrodes revealed high power in the frequency band below 10 Hz, characteristic of EOG signals (Fig. 3m). Collectively, these results confirmed that the T-iSPV electrode effectively measured diverse electrophysiological signals with a performance comparable to that of conventional electrodes.

T-iSPV electrode for stretchable electrochromic displays

In wearable electronics, visual feedback through displays is essential for conveying information to users. Unlike conventional displays, wearable electronics prioritize portability and, therefore, require low-power, low-voltage operation, often powered by batteries. Electrochromic displays have gained attention as low-power display options that use materials that change their absorption wavelength through redox reactions when a voltage is applied.⁵⁴ The implementation of electrochromic displays requires transparent electrodes with high conductivity to transfer charge to the electrochromic layer while exhibiting visible color changes.⁵⁵ Similar to wearable sensors, wearable displays require softness, flexibility, and stretchability to conform to the skin.⁵⁵ Given its stretchability and transparency, the T-iSPV is an ideal candidate for application in stretchable electrochromic displays (sECDs), demonstrating its potential in next-generation wearable display technologies.

P3HT is a well-known electrochromic material;⁵⁶ however, its lack of stretchability hinders its direct application in sECDs. To address this limitation, we dissolved P3HT in chloroform along with SEBS and spin coated the solution to create a stretchable electrochromic composite (sECC) film with enhanced stretchability (Fig. 4a).⁵⁷ The stretchability of the sECC film was maintained, even when it was transferred onto the SEBS substrate (Fig. 4b). The OM images confirmed that the film could stretch by up to 30% without developing cracks, demonstrating stable elongation (Fig. 4c). To evaluate its suitability for ECD applications, the CV measurement of the sECC film was conducted to observe its redox properties (Fig. 4d). Based on these results, an sECD was successfully fabricated by integrating the sECC film with the T-iSPV.

Fig. 4 T-iSPV electrode for stretchable electrochromic displays. (a) Chemical structures of SEBS and P3HT, which compose the stretchable electrochromic composite film. (b) Photograph of the stretchable electrochromic composite (sECC) film transferred onto an SEBS substrate while being stretched. (c) Optical microscope (OM) images of the sECC film before (top) and after stretching (bottom). (d) Cyclic voltammetry data of the sECC film transferred onto a SEBS Au electrode. (e) Schematic illustration showing the device structure of the stretchable electrochromic display (sECD). (f) Absorbance change of the sECD across different wavelengths when voltages from −2 V to 2 V were applied. (g) Photographs showing the color change of the sECD under different applied voltages. (h) Photographs of the sECD attached to the skin under deformation: contraction (top left) and stretching (bottom left), with applied voltages of −2 V and 2 V (top right), and a magnified image (bottom right). (i) Photograph of the sECD stretched up to 30% while applying −2 V (top) and 2 V (bottom). (j) Changes in normalized transmittance over time. (k) Photograph of the sECD operating with PBS drop-casted onto its surface. (l) Photograph of the sECD being stretched underwater while 2 V was applied.

The sECD device consists of top and bottom transparent electrodes for electrical energy transfer, an electrochromic layer, and an electrolyte responsible for ion transport (Fig. 4e and Fig. S17, ESI†). The solid electrolyte consisted of a mixture of poly(vinylidene fluoride-co-hexafluoropropylene) (PVDF-HFP) and the ionic liquid 1-ethyl-3-methylimidazolium bis(trifluoromethyl sulfonyl)imide (EMIM:TFSI). The fabrication of the sECD involved transferring a P3HT composite film onto the T-iSPV electrode, followed by spin-coating the liquid-state ionic electrolyte and covering it with the top T-iSPV electrode. The electrolyte was then solidified to complete the device. The change in transmittance across varying wavelengths when applying a positive voltage (2 V) and a negative voltage (−2 V) to sECD was measured. The sECD exhibited a gradual increase in transparency as the applied voltage increased from −2 V to 2 V, with a corresponding decrease in absorbance in the 550 to 650 nm wavelength range (Fig. 4f). In its default state, without any applied voltage, the sECD displayed a violet hue, became transparent when a positive voltage was applied, and reverted to its original color when a negative voltage was applied (Fig. 4g). Due to the spin-coated, ultrathin layers of the sECC and ion gel combined with the T-iSPV electrode, the sECD maintained conformal contact with the skin and continued to function reliably even under mechanical deformation (Fig. 4h). Moreover, because all the layers of the sECD are stretchable, the device preserved its operation even when subjected to tensile strains of up to 30% (Fig. 4i). The device also demonstrated stable performance with no degradation after being stored in an ambient environment for one hour (Fig. 4j). With proper encapsulation, the sECD maintained its normal functionality even underwater and continued to operate smoothly while being stretched in submerged conditions (Fig. 4k and l). The combination of stretchability, ultrathin form factor, and water resistance makes the sECD highly promising for diverse applications in wearable bioelectronics.

Integrated bioelectronic system based on T-iSVP

Vacuum-deposited electrodes offer high electrical conductivity and fine patterning precision, making them advantageous for biosignal monitoring and electrochromic display operation. However, their intrinsic brittleness and crack-based stretchability often result in a significant increase in resistance under mechanical strain, limiting their direct use in deformable wearable systems. To address this, we applied liquid metal at the connector interfaces to ensure stable electrical contact under deformation, and encapsulated the system with medical tape to enhance mechanical robustness and conformal attachment to the skin.

In the context of wearable electronics, materials must offer not only electrical functionality but also softness, stretchability, and optical transparency for user comfort and seamless integration. The T-iSPV platform satisfies these requirements, enabling its application in both electrophysiological sensing and visual feedback components. To demonstrate its multifunctionality, we developed a fully integrated wearable system by combining T-iSPV-based sensors with sECDs. As shown in Fig. 5a, a single T-iSPV film was used as an ECG electrode and attached directly to the skin using medical tape, forming part of a transparent, skin-conformal bioelectronic interface.

Fig. 5 Integrated bioelectronic system based on T-iSVP. (a) Schematic illustration showing the attachment position, component layout, and operational sequence of the TIBS system. (b) Photograph of the experimental setup used to connect and operate the TIBS system. (c) Block diagram illustrating the system configuration of TIBS, including ECG acquisition, tactile input, and electrochromic display control. (d) Sequential demonstration of the TIBS system operation when attached to the skin, along with the corresponding sECD states and measured signals. The graph displays the state of each sECD (1 = on, 0 = off), the voltage response from the tactile sensing circuit, and the measured heart rate (BPM): (i) Standby mode with the system powered off. (ii) System activation triggered by tactile input, indicated by the first sECD turning on. (iii) Real-time ECG monitoring with the heart rate visually displayed via the first sECD. (iv) Visual alert activated through the second sECD when the heart rate drops below a predefined threshold. (v) Alert deactivated as the heart rate returns to normal, with the second sECD toggling off. [Illustration in (a) was created with BioRender.com].

On the opposite side of the same substrate, two sECDs and one tactile sensor were mounted, forming a transparent integrated bioelectronic display system (TIBS).

In the operating sequence, when the user touches the tactile sensor with a finger, the first sECD is toggled on, visually indicating system activation. Subsequently, ECG monitoring begins, and the heart rate is measured. If the heart rate drops below a preset threshold, the second sECD toggles on to provide a visual alert.

To implement the tactile sensor, two T-iSPV electrodes were arranged with a narrow gap between them and connected in parallel with a reference resistor. When touched by a user, the impedance of skin bridges the gap, resulting in a change in the equivalent resistance of the tactile sensor. Compared to connecting a single T-iSPV electrode in series with the reference resistor, using two electrodes in parallel provided more stable and reliable detection (Fig. S18, ESI†).

As shown in Fig. 5c, the tactile sensor and reference resistor were connected to the Arduino to detect resistance changes. Both sECDs were connected to digital pins on the Arduino and toggled between high (5V) and low (GND) voltage states. Reversing the polarity of these signals enabled the color switching of the sECDs. The ECG electrode transmitted bioelectrical signals to a DAQ system, which processed the data to determine the heart rate. The integrated TIBS system then used this information to control the sECDs, providing a visual representation of the heart rate (Fig. S19 and S20, ESI†).

As shown in Fig. 5d and Video S1 (ESI†), the TIBS system demonstrated smooth operation while attached to the skin, following a clear workflow with visual feedback. (i) In the initial state, the system remains off in standby mode. (ii) When the user touches the tactile sensor, the Arduino detects the resistance change and activates the system, turning on the first sECD to indicate operation. (iii) The system begins measuring the heart rate via the ECG electrode and displays the information on the sECD. (iv) If the heart rate falls below a defined threshold, the second sECD is toggled on to visually alert the user. (v) When the heart rate returns to a normal range, the sECD toggles back off. This comprehensive demonstration of TIBS highlights the seamless integration of ultrathin, stretchable, and transparent components, laying a strong foundation for future wearable bioelectronic systems with multifunctional capabilities.

Conclusions

In this study, we developed an ultrathin, stretchable, and transparent mesh electrode by patterning an Au mesh electrode onto an ultrathin SEBS substrate. To enhance its transparency and stretchability, a lift-off process and optimized pitch, width, and thermal evaporation rate were applied. The T-iSPV electrode demonstrated the ability to detect various biosignals, such as ECG and tactile sensing, while also serving as a transparent electrode for sECD. The T-iSPV electrode holds great promise for a broad range of applications in wearable electronics.

Author contributions

J. J., S. Y., and H. J. contributed equally to this work. J. J., S. Y. and H. J. contributed to conceptualization, data curation, formal analysis, investigation, methodology, visualization, validation and writing original draft. J. Y, J. K., H. C., D. S. (Duhwan Seong) contributed to data curation, resources and software. M. S. and D. S. (Donghee Son) contributed to validation. D. S. (Donghee Son) led the funding acquisition, project administration, supervision, validation, writing – review and editing.

Conflicts of interest

There are no conflicts to declare.

Data availability

The data that support the findings of this study are available from the corresponding author upon reasonable request.

Acknowledgements

This research was supported by the National R&D Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Science and ICT (No. RS-2020-NR047143 and RS-2025-02303342). Also this work was supported by the Institute of Information & Communications Technology Planning & Evaluation (IITP)-ICT Creative Consilience Program grant funded by the Korea government (MSIT) (IITP-2025-RS-2020-II201821). This work was also supported by the Institute for Basic Science (IBS-R015-D1 and IBS-R015-D2). We thank Asahi Kasei for providing SEBS.

Notes and references

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Footnotes

† Electronic supplementary information (ESI) available. See DOI: https://doi.org/10.1039/d5mh00474h

‡ These authors contributed equally.

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