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A universal interface for plug-and-play assembly of stretchable devices


Stretchable hybrid devices have enabled high-fidelity implantable1,2,3 and on-skin4,5,6 monitoring of physiological signals. These devices typically contain soft modules that match the mechanical requirements in humans7,8 and soft robots9,10, rigid modules containing Si-based microelectronics11,12 and protective encapsulation modules13,14. To make such a system mechanically compliant, the interconnects between the modules need to tolerate stress concentration that may limit their stretching and ultimately cause debonding failure15,16,17. Here, we report a universal interface that can reliably connect soft, rigid and encapsulation modules together to form robust and highly stretchable devices in a plug-and-play manner. The interface, consisting of interpenetrating polymer and metal nanostructures, connects modules by simply pressing without using pastes. Its formation is depicted by a biphasic network growth model. Soft–soft modules joined by this interface achieved 600% and 180% mechanical and electrical stretchability, respectively. Soft and rigid modules can also be electrically connected using the above interface. Encapsulation on soft modules with this interface is strongly adhesive with an interfacial toughness of 0.24 N mm−1. As a proof of concept, we use this interface to assemble stretchable devices for in vivo neuromodulation and on-skin electromyography, with high signal quality and mechanical resistance. We expect such a plug-and-play interface to simplify and accelerate the development of on-skin and implantable stretchable devices.

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Fig. 1: BIND connection for stretchable hybrid device.
Fig. 2: Structural analysis of BIND interface and connection.
Fig. 3: In vivo neuromodulation stretchable device assembled through plug-and-play BIND connections.
Fig. 4: A 21-channel on-skin EMG electrode array assembled using plug-and-play BIND connections.

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Data availability

The data supporting the findings of this study are available within the article and its Supplementary Information. Other raw data are available from the corresponding authors upon reasonable request. Source data are provided with this paper.


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G.L. and Zhiyuan Liu acknowledge support from the National Key Research and Development Program of China (grant no. 2021YFF0501601) and the National Natural Science Foundation of China (grant no. 81927804 and U1913601) and NSFC-Shenzhen Robotics Basic Research Center Program (grant no. U2013207). This project is supported by the National Research Foundation, Singapore (NRF) under NRF’s Medium Sized Centre: Singapore Hybrid-Integrated Next-Generation μ-Electronics (SHINE) Centre funding programme. Any opinions, findings and conclusions or recommendations expressed in this material are those of the author(s) and do not reflect the views of NRF. We thank A. L. Chun for critically reading and editing the manuscript. We thank W. Sun and F. Zhao from Bruker Nano Surface Division for AFM characterization.

Author information

Authors and Affiliations



Y.J. and Zhiyuan Liu conceived and coordinated the project and experiments. S.J. assisted the concept development, sample preparation and manuscript writing. J.S. prepared EMG electrodes and collected data in the EMG experiment. G.Z. and H.G. conducted the molecular dynamics in biphasic network growth model. T.S. conducted the Auger characterization and analysis. C.W. assisted the finite element analysis. W.L., H.J. and Zhihua Liu assisted flexible PCB design. M.Y. and Z. Lu assisted EMG electrode design. H.Z. assisted the EMG experiment. G.L. directed, and J.H., J.S. and Y.L. conducted the in vivo animal experiments except bladder. W.Y.X.P. and S.-C.Y. performed the in vivo bladder experiments. J.X., S.W., T.L. and X.Y. assisted BIND interface fabrication. Y.J. prepared all other samples, conducted all other experiments and wrote the manuscript. Z.B. and X.C. directed the project. All authors read and revised the manuscript.

Corresponding authors

Correspondence to Zhiyuan Liu, Zhenan Bao or Xiaodong Chen.

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Competing interests

Y.J., Zhiyuan Liu and X.C. are inventors of the international PCT patent (no. PCT/SG2022/050607, priority date 26 August 2021, filing date 25 August 2022, pending) filed by Nanyang Technological University, which covers the BIND interface and BIND connection reported in this paper.

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Nature thanks Jadranka Travas-Sejdic, Yihui Zhang and Rebecca Kramer-Bottiglio for their contribution to the peer review of this work.

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Extended data figures and tables

Extended Data Fig. 1 Electro-mechanical performance of soft modules connected by conventional pastes or BIND interfaces.

a-d, Schematic (a) and photographs (b-d) of two soft PDMS/Au modules connected using commercial ACF. Stress concentration at connection region results in electrical failure at 50% strain (b, c) and mechanical failure at 76% strain (d). e-h, Schematic (e) and photographs (f-h) of two BIND interfaces connected without any pastes display conductivity up to 180% strain (f, g) and remain connected even at > 600% strain (h). i, Unlike connections using commercial pastes, BIND connection remained conductive (as shown by the low resistance change) even at 180% strain (left). Magnified graph of the low strain region (0–50%) on the right shows BIND connection experienced < 4 times change in relative resistance at 50% strain.

Extended Data Fig. 2 Electro-mechanical performance of BIND interface depends on evaporation rate and thickness.

a, b, Graph (a) and contour map (b) show the electrical stretchability of a single BIND interface as a function of evaporation rate and thickness. At the lowest evaporation rate (0.1–0.2 Å/s) and thickness (45 nm), a non-conductive interface is obtained. As evaporation rate or thickness increases, electrical stretchability peaks before decreasing to ~40%. c, d, Graph (c) and contour map (d) show the electrical stretchability of soft-soft BIND connection as a function of evaporation rate and thickness is similar to the single BIND interface in a and b. Only a limited combination of evaporation rate (0.5–1.0 Å/s) and thickness (45–60 nm) forms the BIND connection. Lower rates and thicknesses result in non-conductive connection while higher ones form non-adhesive interfaces. e, f, Graph (e) and contour map (f) show the mechanical stretchability of soft-soft BIND connection decreasing monotonically with increasing evaporation rate and thickness. Error bars in a, c, e are s.d. from 3–5 samples.

Extended Data Fig. 3 Soft-soft BIND connection is easy to use and has anti-tearing properties.

a-c, Pressing time (a), applied pressure (b) and peeling direction (c) do not influence the adhesion strength of BIND connection very much, indicating that BIND interfaces are easy to use. Simply pressing for 1 s at a low pressure of 0.001 MPa (i.e., a normal finger press) is enough to connect BIND interfaces together. Consistent with the study in Extended Data Fig. 2, non-adhesive connections (120 nm thick Au layer) cannot adhere regardless of pressure. In all other characterization, lap shear test was employed because it’s more common in practical application. d-g, Unlike ACF connection which breaks easily, BIND connection is resistant to cuts. Lap shear test on ACF and BIND connections that were initially cut (~1 mm) in the middle shows BIND connection withstood tearing forces >1.4 N (d) and remained stretchable even at > 500% strain (e). Photos show ACF connection with an initial cut broke easily at the cut location at 50% strain (f), while the initial cut in a BIND connection could still be stretched at 500% strain (g). Error bars in a-c are s.d. from 3-4 samples. Error bars in e are s.d. from 3 samples.

Extended Data Fig. 4 Advantages of BIND connections.

a, b, Schematic (a) and photo (b) showing two patterned BIND interfaces (on module A and B) are joined face-to-face via BIND connection. Inset in b: The BIND connection is electrically conductive to light an LED. c, BIND connections expanded to other conductive materials such as silver/silver and silver/gold connections also display robust electrical and mechanical stretchability. d, Photo of a soft module connected to a flexible PI module patterned with 6 electrode channels, via soft-rigid BIND connection. e, Soft-rigid BIND connection involving rigid or flexible substrates like PI, PET, glass, and metal show higher electrical (~200%) and mechanical (~800%) stretchability than conventional connection via various commercial pastes. f, Interfacial toughness of SEBS encapsulation layer on a BIND interface (0.24 N/mm) is much larger than various other types of encapsulation layer bonded on a conventional PDMS/Au interface. All encapsulation layers are ~100 μm thick. g, SEM image of BIND encapsulation (~300 nm) on a pair of electrodes, exposing two pads for signal collection (SEM shown in magnified view). Scale bar in g: 100 µm. h-j, Resolution of single BIND interface, BIND connection, and BIND encapsulation can achieve 100 µm. The electrical stretchability of both BIND interface (h) and BIND connection (i) decreases with reduced line width, while the mechanical stretchability was kept relatively stable. The overall width was kept as 5 mm. Inset: SEM image of BIND encapsulation with exposed area of 100 × 100 µm2. Scale bar: 50 µm. Error bars are s.d. from 3-4 samples.

Extended Data Fig. 5 Surface AFM mapping shows decreasing ratio of exposed polymer to metal phases, on non-conductive, BIND and non-adhesive interfaces.

a-c, AFM modulus mapping show the polymer/metal ratio on the surface decreases from non-conductive interface (a), to BIND interface (b), to non-adhesive interface (c), which is consistent with the adhesion mapping (Fig. 2a). Due to the large modulus difference, polymer and metal phase can be easily distinguished and are labeled in blue and yellow color, respectively. d-f, AFM height mapping shows similar height variation on non-conductive interface (d) and BIND interface (e), because their height variation comes from the half-immersed gold nanoparticles. For non-adhesive interface (f), the height variation comes from the stacking gold nanoparticles on top of the polymer, so the variation range is slightly different from the non-conductive and BIND interfaces. Scale bar: 100 nm.

Extended Data Fig. 6 Cross-sectional AFM mapping reveals decreasing penetrating Au nanoparticles inside non-conductive, BIND and non-adhesive interfaces.

a, b, AFM height (a) and adhesion (b) mapping of a BIND connection show polymer and metal phases have an interpenetrating structure. Scale bar: 100 nm. c-k, AFM height, adhesion, and current mapping show that the non-conductive connection (c-e) has more interpenetrating Au nanoparticles than the BIND connection (f-h), while there is nearly no interpenetrating Au inside non-adhesive interface (i-k). Non-conductive connection and BIND connection were formed by pressing two interfaces together, while non-adhesive connection was formed by adhering epoxy to non-adhesive interface. White dash line in i shows the epoxy boundary, beyond which the epoxy region was removed for clearer view in j-k. Scale bar in c-k: 400 nm.

Extended Data Fig. 7 In vivo neuromodulation BIND device is compatible with tissues of different sizes, with mechanical robustness against touching and pulling.

a-e, BIND electrode conformally wraps around the common peroneal nerve (a), sciatic nerve (b) and peroneus longus muscle (c) without suturing. Conformal contact is also achieved when placed on the cerebral cortex (d), or sutured onto a bladder wall (e). The bladder experiment can be further improved in the future, to wrap the BIND electrode around the bladder to avoid suture. f-i, Mechanical interference such as touching (f), cathode pulling (g) and anode pulling (h) during sciatic nerve stimulation does not affect the stimulation performance very much. Here, the ultrathin part of the BIND electrode is wrapped around a sciatic nerve while the thick wiring part relays signals via a BIND connection. Stimulated effectiveness was measured by the moving distance of rat ankles upon applied interference (i). Error bars are s.d. from 7–10 movements. j-k, 3D map shows the impedance of BIND interface is typical of a metal-based stretchable electrode (j). Within the frequency range of interest (e.g., ECoG 10-50 Hz, EMG 10-500 Hz), impedance remained low and nearly unchanged up to ~70% strain (k). Interfacial impedance of ultrathin electrode can be improved in the future by various methods (e.g., coating the interface with low impedance materials such as PEDOT:PSS, or increasing the surface area). Because BIND connections construct electrical pathway by Ohmic contact between Au nanoparticles, its influence on impedance is the same as its electrical resistances in DC tests.

Extended Data Fig. 8 The in vivo neuromodulation BIND electrode has various application, including subcutaneous CMAP, cortex ECoG, and bladder, demonstrating its usage universality.

a-f, For simultaneous nerve stimulation and CMAP recording, one BIND electrode is wrapped around the common peroneal nerve and another is wrapped around the peroneus longus muscle (a). Stimulation current was applied to the common peroneal nerve with a parallel resistance of 250 Ω. (b). P-P voltage of CMAP is triggered when stimulation current reaches a threshold, then increases and plateaus (c). Spectroscopy of recorded CMAP shows obvious power density in the typical CMAP frequency range (10–500 Hz) (d). Ankle movement was measured by the distance, d, of rat ankle before (e) and during (f) stimulation. g, h, 4-channel BIND electrode on cerebral cortex for ECoG recording, consisting of ultrathin electrode module, and thick wiring module (g). ECoG signals of healthy and epilepsy rats differ in both amplitude and frequency (h). i-k, For bladder stimulation, two BIND electrodes were sutured onto the bladder wall to transmit electrical pulses, where the bladder contraction is recorded by external pressure sensor (i, j). For bladder recording, one BIND electrode is wrapped around the bladder. Normal saline is injected into the bladder through the catheter, and the corresponding expansion is detected by BIND electrode (k).

Extended Data Fig. 9 The 21-channel EMG electrode assembled via BIND connection shows high signal fidelity and resistance against mechanical interferences, including connection pressing and stretching (50%).

a-b, EMG signals collected from ACF-connected control electrode (a) and BIND electrode (b) show that, pressing/releasing the connection between soft wiring and customized PCB induced noisy signals in control electrode, especially with actions. EMG signals from BIND electrode were less noisy and clearly distinguishable during and after pressure, indicating its resistance to pressure. c-g, Besides pressing, BIND connection also exhibits resistance to 50% stretching interference. One side of the BIND connection between soft wiring and customized PCB was fixed, and the other side was manually stretched to the 50% mark (c). Both electrodes showed noisy signals when stretched but the larger SNR in BIND electrode indicates that it can detect signals under strain (d, e). Releasing strain induced larger noise in control electrode but both electrodes recovered their ability to detect EMG signals after strain is released (f, g). All signals were subjected to a 50 Hz notch filter and a 30 Hz high-pass filter. Error bars in d, e are s.d. from 21 channels.

Extended Data Fig. 10 21-channel EMG electrode assembled via BIND connections maps various hand and finger gestures in air and underwater, with mechanical resistance to interference.

a, BIND electrode detects EMG signals of hand gestures (clench, open, raise and bend), finger movements (stretching of individual fingers) and maximum voluntary contraction (MVC, measured by grip dynamometer) in air. b-d, Even underwater, BIND electrode detected EMG signals with high SNR from different hand and finger gestures (b) and tolerated pressure (c) and 50% strain (d) at connection points. High quality signals with little noise are seen before, during and after pressure or 50% strain. Here, the entire BIND electrode, including two BIND connections, was immersed in water. Signals in a, b were subjected to a 50 Hz notch filter and a 10 Hz high pass filter. Signals in c, d were subjected to a 50 Hz notch filter and a 30 Hz high pass filter.

Supplementary information

Supplementary Information

This file contains Supplementary Notes 1–4, Figs. 1–12, legends for Videos 1–8 and References.

Supplementary Video 1

Robust wiring against mechanical interference. See main Supplementary Information PDF for a full description.

Supplementary Video 2

Simultaneous in vivo CMAP stimulation and recording. See main Supplementary Information PDF for a full description.

Supplementary Video 3

ECoG recording in healthy and epileptic rats. See main Supplementary Information PDF for a full description.

Supplementary Video 4

Plug-and-play assembly of the 21-channel EMG electrode array. See main Supplementary Information PDF for a full description.

Supplementary Video 5

Mechanical interference in 21-channel on-skin EMG detection. See main Supplementary Information PDF for a full description.

Supplementary Video 6

Gesture detection through a 21-channel BIND electrode (in air). See main Supplementary Information PDF for a full description.

Supplementary Video 7

Underwater performance of a 21-channel EMG BIND electrode. See main Supplementary Information PDF for a full description.

Supplementary Video 8

Plug-and-play assembly of a customized multifunctional circuit. See main Supplementary Information PDF for a full description.

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Jiang, Y., Ji, S., Sun, J. et al. A universal interface for plug-and-play assembly of stretchable devices. Nature 614, 456–462 (2023).

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