Permeable stretchable bioelectronic devices and methods of use and making same

Abstract

A permeable stretchable bioelectronic device with wafer-scale patterning capabilities, comprising a stretchable permeable substrate (comprising a biocompatible elastomeric fiber or a biocompatible hydrogel) and one or more electronic components (each independently comprising a metal and a liquid metal) disposed on a surface of the substrate, and methods for making the same.

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims the benefit of priority to U.S. Provisional Patent Application No. 63 / 480,603, filed January 19, 2023, the contents of which are hereby incorporated by reference in their entirety for any purpose.

[0002] The present disclosure relates to permeable stretchable bioelectronic devices and methods of using and making same. [Background technology]

[0003] Implantable skin-mounted bioelectronics are essential for real-time and continuous monitoring of physiological signals for future point-of-care diagnostics, therapeutics, rehabilitation, and augmented reality. Over the past few decades, scientists have made significant progress in developing highly stretchable electronics by patterning devices and circuits on thin polymer substrates (e.g., polyethylene naphthalate, polyimide (PI)), and elastomers (e.g., polydimethylsiloxane (PDMS), polystyrene-ethylene-butylene-styrene, and Ecoflex). With appropriate structural and material design, these stretchable devices not only possess electronic functionality comparable to or superior to that of those fabricated on rigid substrates, but also possess remarkable conformability, enabling conformable contact with soft tissues while avoiding interfacial delamination during movement.

[0004] In addition to high stretchability, long-term (prolonged) biocompatibility is crucial for bioelectronics. It has been reported that the long-term attachment of thin-film (e.g., PDMS) devices to skin and tissue surfaces can cause thermophysiological discomfort (e.g., cold, dampness, skin inflammation, and even malignant tumors) during long-term implantation due to insufficient permeability and a smooth surface. If the polymer substrate (e.g., PI) is too stiff compared to soft tissue, long-term implantation can also lead to tissue scarring. Therefore, in recent years, the development of ultrasoft and permeable types of stretchable electronics has been promoted, where electronic devices are fabricated on porous elastomeric substrates or hydrogels (which offer high stretchability and permeability to air, moisture, and even liquids). These porous elastomeric substrates are primarily fabricated by engineering micropores on polymeric thin films or by direct electrospinning processes. In particular, electrospinning has shown advantages in tunability of porosity, elasticity, and thickness and has been widely applied to the fabrication of various biomedical devices. Hydrogels are typically prepared by crosslinking, polymerization, and gelation of precursor solutions.

[0005] However, due to the high surface roughness and porosity of porous elastomer substrates and the wettability of hydrogels, patterning microelectrodes with high resolution on such ultrasoft and transparent substrates has been extremely challenging. Patternable feature sizes demonstrated to date typically range from 100 μm to several millimeters, resulting in electrode densities of 1–10 electrodes / cm. 2 This is due to the patterning resolution being reduced to a few microns, or 100 electrodes / cm. 2 This is far from meeting the demands of advanced bioelectronics for mapping and intervention with high spatial and temporal resolution, which typically require electrode densities in excess of 1000 kJ / cm. Summary of the Invention [Problem to be solved by the invention]

[0006] Therefore, there is a need for improved methods for making bioelectronic devices and products thereof. [Means for solving the problem]

[0007] In this study, photolithography was used to achieve high resolution down to 2 μm and electrodes with a density of 75,000 electrodes / cm. 2 We report for the first time the wafer-scale patterning of transparent and stretchable liquid metal (LM) microelectrodes (μLMEs) on flexible substrates with ultrahigh patterning densities exceeding 1000%. These μLMEs exhibit tissue-like mechanical flexibility, outstanding electrical conductivity even when stretched to 1,000% strain, and long-term biocompatibility in chronic epicutaneous and implantation tests. We demonstrate an ultrasoft and transparent neural interface for mapping and intervening in electrocorticography (ECoG) signals with high spatiotemporal resolution.

[0008] In a first aspect, presented herein is a permeable stretchable bioelectronic device comprising a flexible substrate and one or more electronic components disposed on a surface of the flexible substrate, wherein the flexible substrate comprises a biocompatible elastomeric fiber or a biocompatible hydrogel, and each of the one or more electronic components independently comprises a metal and a liquid metal.

[0009] In certain embodiments, each of the one or more electronic components is independently selected from the group consisting of a transistor, a diode, an electrode, a generator, an inductor, a capacitor, and a resistor.

[0010] In certain embodiments, the biocompatible elastomeric fiber is selected from the group consisting of an elastomeric homopolymer, an elastomeric block copolymer, an elastomeric random copolymer, an elastomeric graft copolymer, an elastomeric brush copolymer, an elastomeric thermoset, and combinations thereof, and the biocompatible hydrogel comprises a hydrophilic polymer selected from the group consisting of poly(acrylic acid), poly(vinyl alcohol), poly(ethylene oxide), poly(ethylene glycol), gelatin, collagen, polyacrylamide, polysaccharides, and combinations thereof, wherein the hydrophilic polymer is optionally crosslinked.

[0011] In certain embodiments, the biocompatible elastomeric fiber comprises a styrene-isoprene-styrene block copolymer, a styrene-polybutadiene-styrene block copolymer, a styrene-butadiene block copolymer, a poly(styrene-block-butadiene-block-styrene) copolymer, a polyisoprene rubber, a butadiene rubber, or a mixture thereof, and the biocompatible hydrogel comprises crosslinked poly(acrylic acid).

[0012] In certain embodiments, the biocompatible elastomeric fiber comprises a styrene-polybutadiene-styrene block copolymer and the biocompatible hydrogel comprises poly(acrylic acid) crosslinked with N,N'-methylenebisacrylamide.

[0013] In certain embodiments, the biocompatible elastomeric fibers have an average diameter of 0.1 μm to 100 μm.

[0014] In one particular embodiment, a flexible substrate comprising biocompatible elastomeric fibers is plasma treated.

[0015] In one particular embodiment, the flexible substrate is 25 to 330 μm thick.

[0016] In certain embodiments, the metal comprises copper, silver, gold, or a mixture thereof.

[0017] In certain embodiments, the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

[0018] In certain embodiments, the liquid metal comprises a eutectic gallium-indium alloy (EGaIn), a gallium-indium-tin alloy (GaInSn), or a mixture thereof.

[0019] In certain embodiments, the one or more electronic components have a density of 100 to 75,000 electronic components / cm 2 are present on the surface of the flexible substrate at a density of

[0020] In certain embodiments, the flexible substrate comprises poly(styrene-block-butadiene-block-styrene) copolymer or poly(acrylic acid) crosslinked with N,N'-methylenebisacrylamide, the one or more electronic components comprise silver and eutectic gallium-indium alloy (EGaIn), and the flexible substrate comprising poly(styrene-block-butadiene-block-styrene) copolymer is plasma treated.

[0021] In certain embodiments, the one or more electronic components include an electrode.

[0022] In a second aspect, there is provided a method of making a permeable stretchable bioelectronic device as described herein, comprising: providing a substrate having a sacrificial layer formed thereon, the sacrificial layer comprising a water-soluble material selected from the group consisting of water-soluble polymers, water-soluble sugars, water-soluble metal salts, and mixtures thereof; depositing a photoresist layer on the sacrificial layer; irradiating the photoresist layer using a patterned mask, thereby forming an irradiated photoresist layer comprising an exposed photoresist layer and an unexposed photoresist layer; contacting the irradiated photoresist layer with a developer, thereby removing exposed or unexposed photoresist layer and forming a patterned photoresist layer including the exposed patterned sacrificial layer; depositing a metal on the exposed and patterned sacrificial layer; removing the irradiated photoresist layer, thereby forming one or more patterned metal structures on the sacrificial layer; depositing a biocompatible elastomeric fiber onto the one or more patterned metal structures, thereby forming a flexible substrate comprising one or more patterned metal structures disposed on a surface of the flexible substrate, or depositing a hydrogel precursor solution onto the one or more patterned metal structures, thereby forming a hydrogel precursor coating, and curing the hydrogel precursor coating, thereby forming a flexible substrate comprising one or more patterned metal structures disposed on a surface of the flexible substrate; contacting the sacrificial layer with an aqueous solvent, thereby at least partially dissolving the sacrificial layer; Separating the substrate and the flexible substrate including one or more patterned metal structures disposed on a surface of the flexible substrate; depositing a liquid metal onto the one or more patterned metal structures, thereby forming a permeable stretchable bioelectronic device; Presented herein is a method comprising:

[0023] In certain embodiments, the biocompatible elastomeric fibers are deposited using an electrospinning process.

[0024] In one particular embodiment, a flexible substrate comprising biocompatible elastomeric fibers is subjected to a plasma treatment for 1 to 30 minutes.

[0025] In one particular embodiment, the photoresist layer comprises a negative photoresist composition, and the developer removes the unexposed photoresist layer.

[0026] In certain embodiments, the metal comprises copper, silver, gold, or a mixture thereof, and the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

[0027] In certain embodiments, the biocompatible elastomeric fibers are deposited using an electrospinning method, where a flexible substrate is plasma treated for 1 to 30 minutes, the metal comprises copper, silver, gold, or a mixture thereof, and the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

[0028] These and other objects and features of the present disclosure will become apparent from the following description of the disclosure when taken in conjunction with the accompanying drawings. [Brief explanation of the drawings]

[0029] [Figure 1A] Figure 1. Schematic representation of the fabrication process for stretchable and transparent liquid metal microelectrodes (μLMEs). This figure shows the schematic representation of the μLME fabrication process, which consists of four major steps: 1) photolithography of Ag on a SiO2 wafer pre-modified with a thin layer of water-soluble dextran, 2) electrospinning of a fibrous poly(styrene-block-butadiene-block-styrene) (SBS) mat onto the Ag micropattern, 3) dissolution of the dextran layer and transfer of the Ag micropattern from the SiO2 wafer to the SBS fibrous mat, and 4) selective wetting of LM on the Ag-coated areas to generate the μLMEs. [Figure 1B]Figure 1 shows the fabrication process of a stretchable and permeable liquid metal microelectrode (μLME). A schematic diagram comparing different dissolution processes of the sacrificial layer using a permeable fiber mat or an impermeable thin film. Because the solvent can easily penetrate the porous fiber network, using a permeable substrate significantly improves the contact area between the sacrificial layer and the solvent. Therefore, the sacrificial layer can be completely dissolved isotropically within a short time, and the Ag micropattern can be completely transferred. In contrast, when using an impermeable thin film substrate, such as polydimethylsiloxane (PDMS), the solvent can only contact the sacrificial layer from the edge of the substrate, resulting in slow and incomplete dissolution of the sacrificial layer and therefore only partial transfer of the Ag micropattern. [Figure 2A] Figure 1 depicts an exemplary structure of a μLME. Digital images showing the transfer of Ag micropatterns from a wafer to an SBS fiber mat. The inset image shows wafer-sized Ag micropatterns patterned on the wafer before transfer. The main image shows the transfer of Ag to an SBS fiber mat after dissolving the sacrificial layer. The ultrasoft, free-standing sample floats on water. [Figure 2B] 1A-1C are diagrams depicting exemplary structures of μLMEs. Digital image of μLMEs attached to a human arm (LM mass loading: 10 mg / cm2). [Figure 2C] 1 shows an exemplary structure of a μLME. 2 shows the contact angle and penetration rate of a solvent onto a sacrificial layer. [Figure 2D] 1A and 1B are diagrams depicting an exemplary structure of a μLME. 1C and 1D show the pattern transfer success rate as a function of plasma treatment duration. [Figure 2E] Figure 1 represents an exemplary structure of a μLME. Figure 2 Contact angle of the LM on the Ag layer after transfer as a function of Ag thickness. [Figure 2F] 1A-1C are scanning electron microscope (SEM) images showing electrodes on an SBS fiber mat before and after selective wetting by LM. [Figure 2G] 1A-1C are diagrams depicting exemplary structures of μLMEs.

[0023] FIG. 1B is a summary of the line width of μLMEs relative to the original line width of the Ag electrode. [Figure 2H]1A-1C are diagrams depicting exemplary structures of μLMEs. 1D-1C are digital images of high-density μLMEs with a high density of 75,500 electrodes / cm. [Figure 2I] 1A-1C are diagrams depicting exemplary structures of μLMEs. 1D-1C are digital images of high-density μLMEs with a high density of 75,500 electrodes / cm. [Figure 3A] Figure 1 shows experimental results for the permeability, conductivity, and stretchability of μLME. Digital image showing the penetration of a drop of colored water from one side of a μLME to the other. [Figure 3B] Figure 10 depicts experimental results related to the permeability, conductivity, and stretchability of μLMEs. Air permeability of μLMEs deposited on SBS mats with various thicknesses, as well as other common substrates for bioelectronics (including PDMS, Ecoflex, and medical patches). [Figure 3C] Figure 10 depicts experimental results related to the permeability, conductivity, and stretchability of μLME. Moisture permeability of μLME deposited on SBS mats with various thicknesses, as well as other common substrates for bioelectronics (including PDMS, Ecoflex, and medical patches). [Figure 3D] 1 shows experimental results related to the permeability, conductivity, and stretchability of μLMEs. Electrical conductivity of μLMEs with various line widths. [Figure 3E] Figure 10 shows experimental results related to the permeability, conductivity, and stretchability of μLMEs. Resistance of μLMEs with various line widths at different tensile strains. [Figure 3F] Figure 1 shows experimental results related to the permeability, conductivity, and stretchability of μLME. Cyclic electrical stability of μLME (line width 50 μm) under large tensile strain of 1,000%. [Figure 4A] 1A-1C are schematic illustrations showing the application of a μLME-based electrocorticography (ECoG) array and recorded cortical subdomains. [Figure 4B]

[0023] Figure 1 depicts an exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. Digital image of a μLME ECoG electrode array (25 μm thick) in close contact with a flexible, curved, and delicate cerebral cortex. The dashed lines represent the boundaries of the μLME array. [Figure 4C]

[0023] Figure 1 depicts an exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. Digital image of a μLME ECoG electrode array (25 μm thick) in close contact with a flexible, curved, and delicate cerebral cortex. The dashed lines represent the boundaries of the μLME array. [Figure 4D] Figure 1 depicts exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. Figure 2 compares the Young's modulus of μLMEs with previously reported materials for ECoG bioelectrodes. [Figure 4E] Figure 1 illustrates an exemplary high-density μLME-based bioelectronics for neuroelectrophysiology interfaces. In vitro electrochemical impedance spectroscopy (EIS) characterization of μLME and Au / PI electrodes under different strain rates. [Figure 4F] Figure 1 depicts an exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces.Power spectrum analysis of in vivo neural signals performed in a sleep state in a live rat. [Figure 4G] Figure 1 shows an exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interface. Somatosensory evoked potentials (SEPs) recorded using μLME under electrical stimulation with pulsed voltages of various frequencies. [Figure 4H] 1A-1D depict exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. SEP comparison of contralateral and ipsilateral somatosensory cortex under electrical stimulation with pulse voltages of various intensities. [Figure 4I]1A-1C show exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. Spectrograms of signals from the contralateral somatosensory cortex under 10 trains of electrical stimulation with an intensity of 6 V and a frequency of 1 Hz. [Figure 4J] Figure 1 depicts an exemplary high-density μLME-based bioelectronics for neuro-electrophysiology interfaces. Spatiotemporal characteristics of μLME arrays in response to electrical stimulation of the left and right forelimbs. [Figure 5A] Figure 10: Experimental results on the long-term (prolonged) biocompatibility of μLME with prolonged neural implants. Immunohistological analysis of stained brain slices 2 hours after implantation of the electrode array. [Figure 5B] Figure 1 shows experimental results on the long-term biocompatibility of μLME with prolonged neural implantation. Immunohistochemical staining of microglia (Iba1, green) and nuclei (DAPI, blue) for control, Au / PI electrode, Au / PDMS electrode, and μLME electrode at 2 weeks post-implantation. The left panel shows low-magnification images (scale bar = 1,000 μm), and the right panel shows high-magnification images (scale bar = 200 μm). [Figure 5C] Figure 1 shows experimental results related to the long-term (long-term) biocompatibility of μLME with prolonged neural implantation. Statistical analysis of microglial cell body size (n=6). Significant increases in cell body size were observed in the Au / PI group (p=0.0002) and the Au / PDMS group (p=0.0375) compared to the control (two-way, unpaired Student's t-test). No significant differences were observed between the μLME group and the control group (p=0.4482). [Figure 5D]Figure 1 shows experimental results related to the long-term (long-term) biocompatibility of μLME with prolonged neural implantation. Statistical analysis of total microglial cell intensity (n=6). Compared to the control, a significant increase in Iba1 cells was observed in the Au / PI group (p<0.0001) and the Au / PDMS group (p=0.0274) (two-way, unpaired Student's t-test). No significant difference was observed between the μLME group and the control group (p=0.3912). [Figure 6] Digital images depicting the process flow for wafer-scale patterning of transparent, intrinsically stretchable electronics using liquid metal microelectrodes (μLME). The process involves (A) micropatterning of Ag electrodes onto a sacrificial (dextran)-coated wafer using conventional photolithography and lift-off procedures, (B) direct electrospinning of a poly(styrene-block-butadiene-block-styrene) (SBS) fiber mat, used as a transparent stamp / receiving substrate, on top of the Ag micropattern, (C) immersion of the sample in a solvent (water) to remove the sacrificial layer, and (D) and (E) peeling and transfer of the Ag micropattern from the wafer to the SBS fiber mat. [Figure 7A] 1A-1D show experimental results on the morphology of the sacrificial layer coated on the wafer. Atomic force microscope (AFM) topographic images. [Figure 7B] Experimental results on the morphology of the sacrificial layer coated on the wafer. Profile showing the thickness of the sacrificial layer. Polyimide (PI) tape was applied to the surface to identify the uncoated area. The white line indicates the boundary of the tape, and the area to the right of the white line was the uncoated area. [Figure 7C] Figure 1 shows experimental results regarding the morphology of the sacrificial layer coated on the wafer. Surface roughness of the sacrificial layer prepared from solutions of various mass fractions. At a mass fraction of 10 wt%, the resulting sacrificial layer exhibited a relatively low roughness of approximately 0.675 nm. [Figure 7D]Figure 1 shows experimental results regarding the morphology of the sacrificial layer coated on the wafer. Surface roughness of the sacrificial layer prepared from solutions of various mass fractions. At a mass fraction of 10 wt%, the resulting sacrificial layer exhibited a relatively low roughness of approximately 0.675 nm. [Figure 8] 1 shows experimental results for the thickness of electrospun SBS mats as a function of electrospinning duration. By adjusting the electrospinning duration, the thickness of the SBS mats can be well controlled in the range of 25 μm to over 300 μm. [Figure 9] Figure 1 shows experimental results of the topological height of a dextran layer as a function of dissolution time in water, representing the intrinsic dissolution rate of dextran. Dextran spin-coated onto a wafer was directly immersed in water without a fiber mat covering. The dextran layer (approximately 10 nm) completely dissolved within 10 seconds, indicating a fast dissolution rate of approximately 1 nm / s. [Figure 10] Digital images showing water contact angles on an SBS mat before (left) and after (right) plasma treatment. After plasma treatment, the SBS surface became superhydrophilic, with a water contact angle of approximately 0°. [Figure 11] Figure 1 shows experimental results for the bottom wetting time of the solvent (water) as a function of plasma treatment duration during the sacrificial layer release period. For a plasma time of 20 minutes, the water penetration time to wet the bottom surface of the mat decreased to 3 seconds. [Figure 12A] Figure 1 shows digital images of the transfer of the Ag micropattern from the wafer to the SBS mat.Figure 2 shows digital images of the wafer and SBS mat after Ag micropattern transfer. [Figure 12B] Figure 1 shows digital images depicting the transfer of Ag micropatterns from the wafer to the SBS mat.Figure 2 shows digital images depicting the pattern transfer success rate as a function of plasma treatment of the SBS mat. [Figure 13]This figure shows digital images illustrating low-efficiency transfer using an impermeable stamp / receiving substrate. Here, the impermeable substrate was prepared by spin-coating an SBS solution, which formed a transparent thin film on top of the Ag micropattern and served as a control. After plasma treatment, the water contact angle actually decreased from 97° to 32°. However, even after immersing the wafer in water for one week, the Ag micropattern still could not be transferred due to insufficient dissolution of the sacrificial layer and strong adhesion between the SBS film and the underlying Ag micropattern. [Figure 14] SEM and elemental mapping images of μLMEs with various line widths are shown. Due to its poor affinity for SBS and high reactivity with Ag, LM selectively dewets the SBS surface and wets the Ag-covered areas, forming μLMEs on the ultrasoft SBS fiber mat. The LM wetting process originates from reactive alloying on the thick (>100 nm) Ag layer and subsequent wetting of LM on the Ag-In intermetallic compound. Initially, when the LM wetting dosage is low (i.e., 0.139 mg / cm2), the patterned Ag is selectively and reactively wetted primarily by In. As the LM loading increases, the alloy layer is subsequently wetted by LM, revealing a stronger Ga signal in elemental mapping. [Figure 15A] Figure 1 shows the patterning characteristics of μLME. Optical surface profiling images of μLME with line widths ranging from 2 to 200 μm. [Figure 15B] 1 shows the patterning characteristics of μLME. μLME line array with line width and gap of 10 to 50 μm. [Figure 15C] 1 shows the patterning characteristics of μLME. μLME dot arrays with diameters of 50 to 500 μm. [Figure 15D] 10A-10C show the patterning features of μLME: Optical images showing various μLME-based microelectrodes including source-drain electrodes, interconnects, mesh electrodes, and antenna electrodes. [Figure 16]SEM images depicting various electronic components using μLME, including generators, inductors, capacitors, and resistors with linewidths of 4 μm (top) and 8 μm (bottom). The inset image is elemental mapping of element In, providing evidence for the selective wetting of LM. [Figure 17] Digital images showing LM residue under various mechanical pressures. When the electronic patch was attached to the skin, almost no LM residue was observed at load pressures below 3.2 kPa, and only a small amount of LM residue was detected when the load pressure reached 12.8 kPa. [Figure 18] Experimental results regarding the cellular biocompatibility of μLME are shown. (A) Brightfield and fluorescent images of cells cultured in culture medium with a control sample, absorbent gauze, SBS mat, μLME, and 20% DMSO. (B) Quantification of L-929 cell viability in different incubation groups. (C) Absorbance at 450 nm observed by MTT assay after 3 days of incubation for different incubation groups. Brightfield and fluorescent live / dead cell staining images showed regular cell morphology and very few dead cells in all groups except the DMSO group, which showed severe dead cells. In addition, μLME demonstrated a high average cell viability of 98.74% as quantified by live / dead cell staining. The absorbance at 450 nm in the MTT assay (proportional to cell number) for all groups increased significantly with incubation time from days 1 to 3, showing a similar trend, indicating no significant evidence of μLME toxicity. [Figure 19A] 1 shows experimental results for animal testing of μLMEs demonstrating on-skin biocompatibility. Digital images and schematic illustrations show test sites for on-skin application. The inset image was a test μLME with a serpentine structure shape. [Figure 19B] 1 shows experimental results for animal testing of μLMEs demonstrating on-skin biocompatibility. Digital images and schematic illustrations show test sites for on-skin application. The inset image was a test μLME with a serpentine structure shape. [Figure 19C]Figure 1 shows experimental results for animal studies of μLME demonstrating on-cutaneous biocompatibility. Digital images of rabbit skin obtained after wearing various samples during a 72-hour observation period. [Figure 20] Figure 1 shows experimental results for the electrical resistance of μLMEs with various line widths, with the inset image showing the shape and dimensions of the tested samples. [Figure 21] Figure 1 shows experimental results for the electrical conductivity of μLME as a function of LM mass loading. μLME, combined with the underlying Ag layer, has higher electrical conductivity than the original EGaIn. Note that the underlying Ag layer inevitably cracked during the transfer or stretching process due to its brittle nature. By loading the LM, the flowable LM was able to better connect the broken solid metal on the fiber surface and thus maintain a higher electrical conductivity than the as-transferred Ag (which is lower than the theoretical conductivity of Ag). [Figure 22] Experimental results are shown for SEM images showing the morphology of transferred Ag microelectrodes after stretching to 50% strain. (A) Ag microelectrode with a line width of 10 μm. (B) Ag microelectrode with a line width of 50 μm. [Figure 23] Figure 1 shows experimental results for repeated stretch-release tests of μLMEs. (A) μLME with a line width of 10 μm when stretched and released repeatedly between 1000% and 0% strain. (B) μLME with a line width of 200 μm when stretched and released repeatedly between 1500% and 0% strain. [Figure 24] Experimental results show SEM images of a μLME (50 μm) showing the morphological changes of the buckled structure before and after repeated stretch-release tests at 1000% strain. The μLME remains well-connected even after 500 repeated stretch tests at strains ranging from 0 to 1000%. Buckling occurs on the LM due to the formation of a thin solid layer of Ga2O3. Because the oxide layer is much stiffer than the LM, buckling occurs as a result of mechanical competition between the stiffer upper oxide layer and the softer lower substrate. Such a buckled structure contributes to the excellent stretchability and electrical stability of the μLME. [Figure 25]Figure 1 shows experimental results with digital images depicting the experimental setup for neural activity recording and stimulation. (A) Device structure (100 μm thick) of the μLME-based electrocorticography (ECoG) electrode array. (B) ECoG electrode array connected to a printed circuit board (PCB) via a flexible printed circuit (FPC) connector for data acquisition. [Figure 26] Figure 1 shows the experimental results of the stress-strain curves of μLME and Au / PI film. The Au / PI film can only withstand a maximum strain of approximately 30%, while the μLME can withstand a maximum strain of approximately 2,000%. [Figure 27] Figure 1 shows experimental results of digital images depicting the attachment of various electrodes (25 μm thick) to the mouse cortex, including Au / PI film, Au / PDMS, and μLME. Due to its similar mechanical compatibility with brain tissue, μLME (25 μm thick) showed superior compatibility when attached to the cortical surface compared to Au / PI and Au / PDMS films. [Figure 28] Figure 1 shows experimental results of electrochemical impedance spectroscopy (EIS) measurements of μLME and Au / PI films under various strain levels. (A) Phase angle curves of μLME under various strain levels. (B) Phase angle curves of Au / PI films under various strain levels. [Figure 29] Figure 1 shows experimental results for EIS measurements of a multi-channel μLME array as a function of frequency over the range of 1-10,000 Hz. The μLME array exhibited low variation in the electrochemical impedance of each channel. [Figure 30] Figure 1 shows the experimental results of recording neural signals through μLME during sleep. The successful recording of EcoG signals during sleep verified the normal function of μLME for recording cortical activity in vivo. [Figure 31]Figure 1 shows experimental results of the long-term biocompatibility of μLME in skeletal muscle implants. (A) Digital images of electrodes (including μLME, Au / PI, and Au / PDMS) removed after 1 week of implantation. The inset image shows the insertion of the electrode into the skeletal muscle surface. (B) Hematoxylin and eosin (H&E) stained images of skeletal muscle after implantation of various electrodes. [Figure 32] Figures 1A and 1B show experimental results on the long-term stability of μLMEs in air and PBS solution. (A) Digital images showing multi-channel μLME device structures with and without partial encapsulation, where the leads and contact pads are encapsulated in another layer of SBS fiber mat. (B) Electrical resistance of a single-channel μLME as a function of exposure time in air. (C, D) Electrical resistance of a single-channel μLME as a function of exposure time in PBS solution, obtained without encapsulation and with partial encapsulation, respectively. The inset images are schematic illustrations of the test setup. (E, F) Electrochemical impedance of a single-channel μLME as a function of exposure time in PBS solution, obtained without encapsulation and with partial encapsulation, respectively. The inset images are schematic illustrations of the test setup. [Figure 33] Figure 1 shows SEM images of submicron patterned large area LM lines (some of these LM lines were bundled) on an SBS fiber mat. Elemental mapping images confirmed the submicron (down to 250 nm) resolution of this transfer method and the selective wetting of the LM. [Figure 34] This figure shows digital images of the open-cranial surgery procedure for long-term implantation of ECoG electrodes. Three stainless steel bone screws were placed into the skull around the surgical opening, and the dura mater was then removed. Electrodes were placed on the left and right cortical surfaces, respectively. The openings were filled with Kwik-Sil silicone elastomer and then coated with dental cement after the Kwik-Sil cured. [Figure 35] 1 is a table summarizing recently developed flexible / stretchable ECoG devices compared to certain embodiments of the permeable stretchable electrodes described herein. [Figure 36] FIG. 1 shows a table illustrating benchmarks of other patterning techniques for LM using lithographically enabled, implant, additive, and subtractive processes compared to certain embodiments of the method described herein. [Figure 37] 1 is a table illustrating benchmarking of other patterning techniques for transparent stretchable electrodes in terms of resolution, density, advantages, and disadvantages compared to certain embodiments of transparent stretchable electrodes described herein. [Figure 38] 1A-1D are schematic diagrams depicting certain steps for preparing a permeable stretchable bioelectronic device according to certain embodiments herein. (A) depicts an intermediate according to certain embodiments described herein, comprising a flexible substrate (101), one or more patterned metal structures (103), a sacrificial layer (104), and a substrate (105). (B) depicts an intermediate according to certain embodiments described herein, comprising a flexible substrate (101), one or more patterned metal structures (103), and a substrate (105). (C) depicts an intermediate according to certain embodiments described herein, comprising a flexible substrate (101) and one or more patterned metal structures (103). (D) depicts a permeable stretchable bioelectronic device (100) according to certain embodiments described herein, comprising a flexible substrate (101) and one or more electronic components (102). [Figure 39]Figure 1 shows an exemplary fabrication process for liquid metal (LM) micro-iontronics on hydrogel (μLMH). a) Schematic illustration of the μLMH fabrication process, which consists of four major steps: 1) photolithography of Ag on a SiO2 wafer pre-modified with a thin layer of water-soluble dextran; 2) in situ gelation of the hydrogel-receiving substrate onto the Ag micropattern; 3) transfer of the Ag micropattern from the SiO2 wafer to the hydrogel; and 4) selective wetting of LM on the Ag-covered areas to generate μLMHs. b) Schematic illustration showing the tight ion-electron interaction of LM micro-iontronics via metal-carboxylate coordination and ionic interactions. Conventional skin / tissue charge transport between dry-connected metals (e.g., Au, Ag) and LM is weak due to the lack of ionic conductivity and large interfacial impedance. In contrast, wetting ionic gels overcome these drawbacks but suffer from low electrical conductivity. LM iontronics can provide a low-impedance biointerface to skin / tissue due to the tight ion-electron interactions and negligible spreading resistance. c, Digital image showing low-density and high-resolution μLMHs functioning as a transparent and seamless biointerface. The inserted scanning electron microscope (SEM) image illustrates the cross-shaped structure of μLMHs. d, Digital images of high-density μLMHs of various shapes and dimensions, including lines, dots, wires, and interconnects for stretchable circuits. [Figure 40A] Figure 1 shows experimental results for low impedance μLMH resulting from tight ion-electron interactions. Time-of-flight secondary ion mass spectrometry (TOF-SIMS) cation depth (Ga and In ions) 3D profiles showing ionic interactions between the LM and PAA hydrogel in μLMH. [Figure 40B] Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. X-ray photoelectron spectroscopy (XPS) spectra obtained for C1s(h) of PAA / LM(μLMH), Ga3d with In3d and In4d of PAA / LM, respectively. [Figure 40C]Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. X-ray photoelectron spectroscopy (XPS) spectra obtained for C1s(h) of PAA / LM(μLMH), Ga3d with In3d and In4d of PAA / LM, respectively. [Figure 40D] Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. X-ray photoelectron spectroscopy (XPS) spectra obtained for C1s(h) of PAA / LM(μLMH), Ga3d with In3d and In4d of PAA / LM, respectively. [Figure 40E] Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. Fourier transform infrared (FTIR) spectra of PAA with and without LM coating. [Figure 40F] Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. Bode plots of ion-conducting PAA with various current collectors, including pristine stainless steel, Au, and LM. The inset image is a schematic illustration of LM / PAA / LM. [Figure 40G] Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions: Nyquist plot of LM / PAA / LM (ionic conductivity of PAA is about 2.73 S / m) with typical Warburg impedance, indicating the presence of ionic diffusion at the LM / PAA interface. [Figure 40H] Figure 1 shows experimental results for low impedance μLMH resulting from tight ion-electron interactions. Ionic conductivity of PAA with various current collectors. [Figure 40I] Experimental results for low-impedance μLMH resulting from tight ion-electron interactions. Skin impedance at various biological interfaces, including commercial electrodes (Ag / AgCl gel), metal electrodes (LM / PDMS), and wet electrodes using LM and a conventional metal (Au) as the current collector. [Figure 40J]Figure 1 shows experimental results for low-impedance μLMH resulting from tight ion-electron interactions. Electrochemical impedance of various biological interfaces including dry metal (Au), LM electrodes, hydrogel, and μLMH electrodes, respectively. [Figure 41] Experimental results on high-resolution μLMH patterned from tuning the interfacial adhesion. a) Storage and loss moduli of polyacrylic acid (PAA) as a function of gelation duration. b, c) Interfacial adhesion of the transfer stamp (PAA) and pattern transfer rate of Ag microelectrodes as a function of gelation duration. The inset image is a schematic illustration of the test setup for interfacial shear adhesion strength. d) SEM images showing transparent mesh electrodes on polyimide (PI, before transfer) and PAA (after transfer) before and after selective wetting with LM. e) Optical surface profiling images of μLMH as a function of LM mass loading, respectively. [Figure 42A] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMH. Young's modulus and stress-strain curves of μLMH prepared with precursor solutions (PAA dissolved in HO / glycerin) of various mass ratios. [Figure 42B] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMH. Young's modulus and stress-strain curves of μLMH prepared with precursor solutions (PAA dissolved in HO / glycerin) of various mass ratios. [Figure 42C] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMHs. Transmittance as a function of wavelength for μLMHs with various patterning resolutions (line width and gap) ranging from 5 μm to 100 μm. [Figure 42D] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMH. Electrical conductivity of μLMH with various LM mass loadings. [Figure 42E] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMH. Electrical resistance of μLMH (50 mm line width and gap, average 53.7% transmittance at 550 nm) during stretch-relaxation tests from 0% to 1,000% strain. [Figure 42F] Figure 1 shows experimental results on the stretchability, transparency, and electrical conductivity of μLMHs. Digital images of μLMHs with high electrical conductivity that function as wiring and interconnects for analog circuits. [Figure 43A] Figure 1 shows experimental results of transcutaneous electrical stimulation and electrophysiological sensing functions for low-impedance μLMH. Transcutaneous electrical stimulation was performed using dry Au, μLM, hydrogel, and μLMH electrodes, respectively, under various electrical stimulation voltages and frequencies. [Figure 43B] Figure 1 shows experimental results of transcutaneous electrical stimulation and electrophysiological sensing functions for low-impedance μLMH. Transcutaneous electrical stimulation was performed using dry Au, μLM, hydrogel, and μLMH electrodes, respectively, under various electrical stimulation voltages and frequencies. [Figure 43C] Figure 1 shows experimental results for transcutaneous electrical stimulation and electrophysiological sensing functions of low-impedance μLMH. Corresponding electromyographic (EMG) signals of the forearm recorded using various biological interfaces including dry Ag, LM, hydrogel, and μLMH electrodes, respectively, in terms of signal-to-noise ratio (SNR) and response time. [Figure 43D] Figure 1 shows experimental results of transcutaneous electrical stimulation and electrophysiological sensing functions for low-impedance μLMH. Corresponding electromyogram (EMG) signals of the forearm using various biointerfaces including dry Ag, LM, hydrogel, and μLMH electrodes, respectively, in terms of signal-to-noise ratio (SNR) and response time. [Figure 43E] Figure 1 shows experimental results of the low-impedance μLMH for transcutaneous electrical stimulation and electrophysiological sensing functions. EMG signal recordings of various biological interfaces under specific electrical stimulation (1 Hz, 20 V). [Figure 43F] 1 shows experimental results of transcutaneous electrical stimulation and electrophysiological sensing functions on a low-impedance μLMH. Recordings of EMG signals of the μLMH under electrical stimulation with different frequency values. [Figure 43G]Figure 1 shows experimental results of the low-impedance μLMH for transcutaneous electrical stimulation and electrophysiological sensing functions: Spectrograms of EMG signals under 20 consecutive electrical stimulations with an intensity of 20 V and a frequency of 100 Hz. [Figure 44A] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuro-electrophysiological interface with optogenetic mapping. Schematic illustration and digital images showing the experimental setup and stimulation area for optogenetic mapping using a 32-channel μLMH and 473 nm blue light. [Figure 44B] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuro-electrophysiological interface with optogenetic mapping. Schematic illustration and digital images showing the experimental setup and stimulation area for optogenetic mapping using a 32-channel μLMH and 473 nm blue light. [Figure 44C] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuro-electrophysiological interfaces with optogenetic mapping capabilities. Schematic illustration of a multi-layer and multi-channel μLMH-based electrocorticography (ECoG) array for hydrogel bioelectronics. [Figure 44D] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuroelectrophysiological interfaces with optogenetic mapping capabilities. Electrochemical impedance of a 32-channel μLMH as a function of frequency. [Figure 44E] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuro-electrophysiological interface with optogenetic mapping. Spectrogram of ECoG signals from the contralateral somatosensory cortex under continuous light stimulation (473 nm blue light). [Figure 44F]Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuro-electrophysiological interface with optogenetic mapping function. Spatiotemporal characteristics of μLMH array in response to optical stimulation (473 nm blue light). [Figure 44G] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuroelectrophysiological interfaces with optogenetic mapping capabilities. Immunohistological analysis of stained brain slices after μLMH implantation. [Figure 44H] Figure 1 shows experimental results of high-density μLMH-based hydrogel bioelectronics for neuroelectrophysiological interface with optogenetic mapping. Statistical analysis of microglial cell body size (n=6). No significant difference was observed between the μLMH group and the control group (P=0.8469). [Figure 45] Figure 1 shows experimental results for liquid metal (LM) micro-iontronics (μLMH) on hydrogels used as low-impedance biointerfaces. a) Schematic illustration showing the current interface between biology and electronics, mainly including gel electrodes, dry-contact electrodes, and iontronics. b) Equivalent circuit of the electrode / tissue interface. Neural activity can generate a voltage source (Ve). Rspread is the resistance of the extracellular space, which depends on the geometry of the recording site. Re and Ce are, in the simplest case, attributed to the leakage resistance and double-layer capacitance of the electrode itself, i.e., the electrode / tissue interface. [Figure 46] Figure 1 shows experimental results for digital images depicting μLMHs with various shapes, dimensions, and transparencies, including a)-c) dot and line arrays, d) transparent electrodes, and stretchable circuits. [Figure 47] FIG. 1 shows experimental results of μLMH impedance as a function of frequency for PAA hydrogels with different ionic conductivities, using LM as the current collector. [Figure 48] FIG. 1 shows experimental results on the electrochemical impedance of μLMHs with various line widths ranging from 10 μm to 80 μm. [Figure 49] Figure 1 shows a schematic illustration of the modulation of interfacial adhesion during the transfer process.The transfer printing process consists of collecting / picking up ink from a donor substrate and printing / delivering ink to a receiving substrate. [Figure 50] FIG. 1 shows atomic force microscope topographic images depicting the dissolution rate of the sacrificial layer (dextran). [Figure 51A] Figure 1 shows experimental results for the characterization of polyacrylic acid (PAA) hydrogels. Schematic of PAA synthesis. [Figure 51B] 1 shows experimental results for the characterization of polyacrylic acid (PAA) hydrogels: Proton nuclear magnetic resonance (NMR) spectroscopy of AA and PAA. [Figure 51C] 1 shows experimental results for the characterization of polyacrylic acid (PAA) hydrogels: Raman spectra of PAA as a function of crosslinking duration. [Figure 52] FIG. 10 shows optical images depicting the transfer status of Ag patterns as a function of the crosslinking time of the hydrogel stamp. [Figure 53] Experimental results on PAA stamp adhesion: a) Interfacial shear adhesion strength of μLMH when adhered to various substrates including PI, PC, PDMS, and PET films. b) Digital images showing high adhesion of μLMH to various porcine organs including tongue, liver, pig skin, and heart. [Figure 54] Figure 1 shows experimental results for a) thermogravimetric analysis of PAA stamps with various mass fractions of glycerin in the precursor solution, b) digital images showing the anti-dehydration properties of PAA (50 wt% glycerin) compared to PAA hydrogels without glycerin. [Figure 55] FIG. 1 shows scanning electron microscope (SEM) images of μLMHs with various mass loadings of LMs and corresponding elemental mapping images. [Figure 56]FIG. 1 shows experimental results for the electrical resistance of μLMH (10 μm line width and gap, average 79.8% transmittance at 550 nm) during stretch-relaxation testing from 0% strain to 1,000% strain. [Figure 57] FIG. 1 shows a schematic illustration representing electrocorticogram (ECoG) cortical subdomains recorded in the rat cerebral cortex using μLMH arrays. [Figure 58] FIG. 1 shows ECoG signals from a sleeping rat recorded using a 32-channel μLMH array. [Figure 59] FIG. 1 shows the amplitude of ECoG signals from living rats in a sleep state (without light stimulation) and with light stimulation (473 nm blue light). [Figure 60] 1 is a table showing benchmarks compared to other patterning technologies in iontronics in terms of resolution, density, advantages, and disadvantages. DETAILED DESCRIPTION OF THE INVENTION

[0030] definition Throughout this disclosure, unless the context otherwise requires, the word "comprise" or derivatives thereof, such as "comprises" or "comprising," will be understood to refer to the inclusion of a stated integer or group of integers, and not the exclusion of other integers or groups of integers. Note also that in this disclosure, and particularly in the claims and / or paragraphs, terms such as "comprises," "comprised," "comprising," and the like may have the meaning ascribed to them in U.S. patent law, e.g., "includes," "included," "including," and the like, and terms such as "consisting essentially of" and "consists essentially of" have the meaning ascribed to them in U.S. patent law, e.g., allowing for elements not expressly recited but excluding elements found in the prior art or that affect a basic or novel characteristic of the invention.

[0031] Furthermore, throughout this disclosure and the claims, unless the context otherwise requires, the word "include" or derivatives thereof, such as "includes," "including," etc., will be understood to imply the inclusion of a stated integer or group of integers, and not the exclusion of other integers or groups of integers.

[0032] The use of the singular herein includes the plural (and vice versa) unless specifically stated otherwise. In addition, when the term "about" is used before a quantitative numerical value, the present teachings also include the specific quantitative numerical value itself unless specifically stated otherwise. As used herein, the term "about" refers to a variation of ±10%, ±7%, ±5%, ±3%, ±1%, or ±0% from the nominal value, unless otherwise indicated or inferred.

[0033] As used herein, a "polymeric compound" (or "polymer") refers to a molecule comprising a plurality of one or more repeating units joined by covalent chemical bonds. A polymeric compound has the general formula I: *-(-(Ma) x -(Mb) y -) z * General formula I It can be expressed as:

[0034] where Ma and Mb each represent a repeating unit or monomer. A polymeric compound may have only one type of repeating unit, as well as two or more different types of repeating units. If a polymeric compound has only one type of repeating unit, it may be called a homopolymer. If a polymeric compound may have two or more different types of repeating units, the term "copolymer" or "copolymer compound" may be used instead. For example, a copolymeric compound may include a repeating unit in which Ma and Mb represent two different repeating units. Unless otherwise specified, the incorporation of repeating units within a copolymer may be head-to-tail, head-to-head, or tail-to-tail. In addition, unless otherwise specified, a copolymer may be a random copolymer, an alternating copolymer, or a block copolymer. For example, general formula I can be used to describe a copolymer of Ma and Mb having x mole fraction of Ma and y mole fraction of Mb in the copolymer, where the repeating comonomers Ma and Mb can be alternating, random, regiorandom, regioregular, or blocky, and there are a maximum of z comonomers. In addition to their composition, polymeric compounds can be further characterized by their degree of polymerization (n) and molar mass (e.g., number average molecular weight (M) and / or weight average molecular weight (Mw) depending on the measurement technique). The polymers described herein can exist in a wide variety of stereochemical configurations (e.g., isotactic, syndiotactic, atactic, or combinations thereof).

[0035] As used herein, a "biocompatible elastomeric polymer" refers to an elastomeric polymer that is substantially non-toxic to cells or organisms with acceptable tolerance, including being substantially non-carcinogenic and substantially non-immunogenic in a subject.

[0036] As used herein, a "fiber" is an elongated, thin, thread-like, and / or fibrous structure. A "fiber layer" is any two- or three-dimensional arrangement of fibers (either an ordered arrangement of fibers (e.g., a woven or non-woven mesh) or a random arrangement of fibers (typically a mat of fibers produced by electrospinning)), and can be isotropic or anisotropic.

[0037] As used herein, the terms "flexible," "bendable," or "stretchable" refer to the ability of a material, structure, device, or device component to deform, e.g., into a curved or folded shape that remains intact during bending, folding, or stretching, without undergoing a transformation that results in significant strain (e.g., a strain that characterizes the yield point of the material, structure, device, device component, etc.).

[0038] As used herein, the term "hydrogel" refers to a three-dimensional network of hydrophilic polymers that contains water. Hydrogels can contain, for example, 10%, 20%, 30%, 40%, 50%, 60%, 70%, 80%, 90%, 95%, or even more water on a w / w basis.

[0039] As used herein, the term "hydrogel precursor" refers to a monomer or polymer that is at least partially soluble in aqueous media and has the ability to become crosslinked to form a hydrogel.

[0040] Referring to Figure 38D, the present disclosure provides a permeable stretchable bioelectronic device (100) comprising a flexible substrate (101) and one or more electronic components (102) disposed on a surface of the flexible substrate, wherein the flexible substrate comprises a biocompatible elastomeric fiber or a biocompatible hydrogel, and each of the one or more electronic components independently comprises a metal and a liquid metal. In the case where the flexible substrate comprises a biocompatible elastomeric fiber, the flexible substrate is a fiber layer.

[0041] The type of electronic component is not particularly limited and can be any electronic component that can be fabricated using lithography. Exemplary electronic components include, but are not limited to, transistors, diodes, electrodes, generators, inductors, capacitors, resistors, organic electrochemical transistors (OECTs), coils, lines, dots, source-drain electrodes, interconnects, mesh electrodes, and antennas. In certain embodiments, one or more electronic components include electrodes.

[0042] The physical characteristics of one or more electronic components (e.g., line dimensions (width, height, thickness, etc.), line spacing, dots (e.g., diameter), etc.) may range in size from 0.25 to 500 μm, 0.25 to 400 μm, 0.25 to 300 μm, 0.25 to 200 μm, 0.25 to 100 μm, 0.25 to 75 μm, 0.25 to 50 μm, 0.25 to 25 μm, 0.25 to 20 μm, 0.25 to 15 μm, 0. It can be in the range of 25 to 10 μm, 0.25 to 5 μm, 0.25 to 4 μm, 0.25 to 3 μm, 0.25 to 2 μm, 0.25 to 1 μm, 0.50 to 500 μm, 0.75 to 500 μm, 1 to 500 μm, 2 to 500 μm, 3 to 500 μm, 4 to 500 μm, 5 to 400 μm, 5 to 300 μm, 5 to 200 μm, 5 to 100 μm, 5 to 50 μm, 5 to 25 μm, or 10 to 500 μm.

[0043] Advantageously, the methods described herein can be used to fabricate permeable stretchable bioelectronic devices comprising a high density of electronic components, where the one or more electronic components may be sized to have a density of 1-100,000, 1-90,000, 1-80,000, 1-75,000, 10-75,000, 100-75,000, 1,000-75,000, 10,000-75,000, 20,000-75,000, 30,000-75,000, 40,000-75,000, 50,000-75,000, 55,000-75,000, or 60,000-75,000. 00, 60,000-75,000, 65,000-75,000, 70,000-75,000, 71,000-75,000, 72,000-75,000, 73,000-75,000, 74,000-75,000, 1-2,500, 100-2,500, 500-2,500, 1,000-2,500, 1,500-2,500, or 2,000-2,500 pieces / cm 2 The nanoparticles may be present on the surface of the flexible substrate at a density of

[0044] The biocompatible elastomeric fibers can be elastomeric homopolymers, elastomeric block copolymers, elastomeric random copolymers, elastomeric graft copolymers, elastomeric brush copolymers, elastomeric thermosets, or combinations thereof.

[0045] In certain embodiments, the biocompatible elastomeric fiber comprises a styrene-isoprene-styrene block copolymer, a styrene-polybutadiene-styrene block copolymer, a styrene-butadiene block copolymer, a poly(styrene-block-butadiene-block-styrene) copolymer, polyisoprene rubber, butadiene rubber, polyurethane, thermoplastic polyurethane, polyvinyl alcohol, polycaprolactone, polycaprolactone, or a mixture thereof. In certain embodiments, the biocompatible elastomeric fiber comprises poly(styrene-block-butadiene-block-styrene). In certain embodiments, the biocompatible elastomeric fiber further comprises one or more functional groups selected from the group consisting of hydroxyl, carboxyl, epoxy, and keto.

[0046] The biocompatible elastomeric fibers can have an average diameter of 0.1 μm to 100 μm, 0.1 μm to 50 μm, 0.1 μm to 40 μm, 0.1 μm to 30 μm, 0.1 μm to 20 μm, 0.1 μm to 10 μm, 0.5 μm to 10 μm, 1 μm to 10 μm, 1 μm to 9 μm, 1 μm to 8 μm, 1 μm to 7 μm, 1 μm to 6 μm, 1 μm to 5 μm, 1 μm to 4 μm, or 1 μm to 3 μm. In certain embodiments, the biocompatible elastomeric fibers have an average diameter of about 2 μm.

[0047] In certain embodiments, the biocompatible hydrogel comprises a hydrophilic polymer selected from the group consisting of poly(acrylic acid), poly(vinyl alcohol), poly(ethylene oxide), poly(ethylene glycol), gelatin, collagen, polyacrylamide, polysaccharides, and combinations thereof, and the hydrophilic polymer is optionally crosslinked. In cases where the hydrophilic polymer is crosslinked, the crosslinking can be achieved by covalent conjugation, non-covalent conjugation (e.g., by ion-ion interactions, ion-hydrogen bonding interactions, hydrogen bonding interactions, or a combination thereof), or a combination thereof. Exemplary crosslinkers include methacrylate / acrylate-based crosslinkers (e.g., ethylene glycol dimethacrylate, diethylene glycol diacrylate, diethylene diacrylate, diethylene acrylate, allyl methacrylate, or 1,4-butanediol diacrylate, allyl methacrylate, or 1,4-butanediol dimethacrylate, poly(ethylene glycol) dimethacrylate, poly(ethylene glycol) diacrylate, polyethylene oxide dimethacrylate, polyethylene oxide diacrylate, N,N'-methylenebisacrylate, Crosslinkers include, but are not limited to, N,N'-methylenebis(2-methylacrylamide), methylene dimethacrylate, N,N'-methylenebis(2-methylacrylamide), methylene diacrylate, methylene bis(2-methylacrylate), diethylene glycol diacrylate, hexamethylene diacrylate, oxybis(methylene)bis(2-methylacrylate), and oxybis(ethane-2,1-diyl)bis(2-methylacrylate), as well as metal salt-based crosslinkers (e.g., sodium chloride, calcium chloride, calcium nitrate, etc.). In certain embodiments, the biocompatible hydrogel comprises polyacrylic acid crosslinked with N,N'-methylenebisacrylamide.

[0048] The thickness of the flexible substrate can be in the range of 5 to 1,000 μm, 5 to 750 μm, 5 to 500 μm, 5 to 400 μm, 10 to 400 μm, 15 to 400 μm, 20 to 400 μm, 25 to 400 μm, 25 to 350 μm, 25 to 330 μm, 50 to 350 μm, 100 to 350 μm, 150 to 350 μm, 200 to 350 μm, 250 to 350 μm, 300 to 350 μm, 25 to 300 μm, 25 to 250 μm, 25 to 200 μm, 25 to 150 μm, 25 to 100 μm, 25 to 50 μm, 50 to 300 μm, 100 to 250 μm, or 150 to 200 μm.

[0049] As discussed herein, in instances where the flexible substrate comprises biocompatible elastomeric fibers, plasma treatment of the flexible substrate is advantageous for facilitating the dissolution and removal of sacrificial layers used to fabricate the permeable stretchable bioelectronic devices described herein. Accordingly, in certain embodiments, a flexible substrate comprising biocompatible elastomeric fibers is plasma treated to incorporate hydrophilic moieties (e.g., hydroxyl, carboxyl, epoxy, keto, etc.) into and / or on the biocompatible elastomeric fibers. A flexible substrate comprising biocompatible elastomeric fibers can be plasma treated in an atmosphere containing oxygen and / or nitrogen on the side of the flexible substrate opposite one or more electronic components. In certain embodiments, a flexible substrate comprising biocompatible elastomeric fibers is plasma treated in an atmosphere.

[0050] The metal present in the one or more electronic components is not particularly limited and can be any metal used in the fabrication of electronic components. In certain embodiments, the metal is substantially non-toxic to cells or organisms with acceptable tolerance, including being substantially non-carcinogenic and substantially non-immunogenic in a subject. In certain embodiments, the metal comprises copper, silver, gold, magnesium, molybdenum, iron, or a mixture thereof.

[0051] The choice of liquid metal is not particularly limited, but is preferably a metal that is substantially non-toxic to cells or organisms with acceptable tolerability, including being substantially non-carcinogenic and substantially non-immunogenic in the subject. In certain embodiments, the liquid metal comprises gallium, a gallium alloy, or a mixture thereof. Exemplary liquid metals include, but are not limited to, EGaIn, GaInSn, Ga / Sn, or a mixture thereof.

[0052] The permeable stretchable bioelectronic devices described herein exhibit improvements in air and moisture permeability compared to bioelectronic devices known in the art. In certain embodiments, a permeable stretchable bioelectronic device described herein comprising a flexible SBS substrate having a thickness of about 330 μm exhibits an air permeability of about 78 mm / sec and a moisture permeability of about 835 g / m 2 In certain embodiments, a permeable stretchable bioelectronic device described herein comprising a flexible SBS substrate having a thickness of about 25 μm exhibits an air permeability of about 235 mm / sec and a moisture permeability of about 990 g / m 2 Indicates moisture permeability per day.

[0053] The permeable stretchable bioelectronic devices described herein can exhibit a maximum strain of about 2,000%. In certain embodiments, the permeable stretchable bioelectronic devices exhibit a maximum strain of between 100% and 500%, 500 and 2,000%, 1,000 and 2,000%, 1,500 and 2,000%, 500 and 1,500%, 500 and 1,000%, or 1,000 and 1,500%.

[0054] The permeable stretchable bioelectronic devices described herein can be bioelectronic electrodes, for example, bioelectronic electrodes can be used in medical devices for EMG, ECG, EEG, hydration sensing, muscle monitoring, and impedance measurement.

[0055] The present disclosure provides a method of making the permeable stretchable bioelectronic device described herein, comprising the steps of: (a) providing a substrate having a sacrificial layer formed thereon, wherein the sacrificial layer comprises a water-soluble material selected from the group consisting of a water-soluble polymer, a water-soluble sugar, a water-soluble metal salt, and mixtures thereof; (b) depositing a photoresist layer on the sacrificial layer; (c) irradiating the photoresist layer using a patterned mask, thereby forming an irradiated photoresist layer comprising an exposed photoresist layer and an unexposed photoresist layer; (d) contacting the irradiated photoresist layer with a developer, thereby removing the exposed or unexposed photoresist layer and forming a patterned photoresist layer comprising an exposed, patterned sacrificial layer; (e) depositing metal on the exposed, patterned sacrificial layer; (f) removing the irradiated photoresist layer, thereby forming one or more patterned metal structures on the sacrificial layer; and (g) removing one or more (i) contacting the sacrificial layer with an aqueous solvent, thereby at least partially dissolving the sacrificial layer; (j) separating the substrate and the flexible substrate comprising the one or more patterned metal structures disposed on the surface of the flexible substrate; and (k) depositing a liquid metal on the one or more patterned metal structures, thereby forming a flexible substrate comprising the one or more patterned metal structures disposed on the surface of the flexible substrate.

[0056] Figure 38 illustrates steps (i)-(k) of the method described herein, where in step (i), a sacrificial layer (104) comprising a water-soluble polymer is dissolved, thereby at least partially removing the sacrificial layer (104) from between the substrate (105) and one or more patterned metal structures (103), as illustrated in Figures 38A-38B. In step (j), the substrate (105) is separated from the flexible substrate (101), which includes one or more patterned metal structures (103) disposed on the surface of the flexible substrate, as illustrated in Figures 38B-38C. In step (k), a liquid metal is deposited on the one or more patterned metal structures (103) [shown in Figure 38C (103)], thereby forming a permeable stretchable bioelectronic device, shown in Figure 38D, including the flexible substrate (101) and one or more electronic components (102) disposed on the surface of the flexible substrate (101).

[0057] The substrate can be silicon dioxide, silicon, polyimide, polyethylene terephthalate, polycrystalline silicon, silicon nitride, tetraethyl orthosilicate, amorphous carbon, silicon oxynitride, germanium, silicon germanium, diamond, quartz, or a semiconductor.

[0058] The water-soluble polymer is not particularly limited and can be any polymer that is at least partially soluble in water. Exemplary water-soluble polymers include, but are not limited to, carbohydrates (e.g., amylose, dextran, dextrin, maltose, isomaltose, maltotriose, and stachyose, polyacrylic acid, polymethylacrylic acid, polyvinyl alcohol, polyacrylamide, polyalkylene oxide, and mixtures thereof). In certain embodiments, the water-soluble polymer is dextran.

[0059] The water-soluble sugars can be monosaccharides, disaccharides, trisaccharides, tetrasaccharides, oligosaccharides, or mixtures thereof. Exemplary water-soluble sugars include, but are not limited to, glucose, galactose, fructose, xylose, sucrose, lactose, maltose, isomaltulose, trehalose, sorbitol, and mannitol.

[0060] The water-soluble metal salt may comprise a metal selected from Groups 1, 2, 4, 10, 11, and 12 of the periodic table and one or more anions selected from the group consisting of halide (e.g., chloride, bromide, and iodide), carbonate, nitrate, sulfate, bicarbonate, phosphate, monohydrogen phosphate, dihydrogen phosphate, sulfamate, and acetate. In certain embodiments, the water-soluble metal salt is a calcium salt comprising one or more anions selected from the group consisting of chloride, bromide, carbonate, nitrate, sulfate, bicarbonate, phosphate, and monohydrogen phosphate.

[0061] Deposition of the sacrificial layer can be achieved using any number of techniques, such as spin coating of an aqueous solution containing a water-soluble polymer, printing, print screening, spraying, painting, doctor blading, slot-die coating, or dip coating, etc. In one particular embodiment, the sacrificial layer is deposited on the substrate by spin coating a solution containing a water-soluble material onto the substrate surface.

[0062] The photoresist layer may comprise a negative photoresist composition or a positive photoresist composition. Examples of positive photoresist compositions include, but are not limited to, polymethyl methacrylate, diazoquinone esters, and phenolic novolac resins. Examples of negative photoresist compositions include, but are not limited to, acrylate-based photoresists and epoxy-based photoresists (e.g., SU-8).

[0063] Deposition of the photoresist layer can be achieved using any number of techniques, such as spin coating, printing, print screening, spraying, painting, doctor blading, slot die coating, or dip coating of a solution containing the photoresist composition.

[0064] The step of irradiating the photoresist layer may be carried out using a lithographic method selected from the group consisting of electron beam lithography, optical lithography, ultraviolet lithography, ion beam lithography, X-ray lithography, interference lithography, scanning probe lithography, charged particle lithography, or nanoimprint lithography.

[0065] The biocompatible elastomeric fibers can be deposited onto one or more patterned metal structures using methods known in the art for depositing fibers onto substrate surfaces. In certain embodiments, the biocompatible elastomeric fibers are deposited using electrospinning.

[0066] The hydrogel precursor solution may include one or more hydrogel precursors selected from the group consisting of polymerizable monomers, hydrophilic polymers, a crosslinker, and a solvent. The polymerizable monomers may be acrylates, methacrylates, acrylic acid, methacrylic acid, and combinations thereof. The crosslinker may be ethylene glycol dimethacrylate, diethylene glycol diacrylate, diethylene diacrylate, diethylene acrylate, allyl methacrylate, or 1,4-butanediol diacrylate, allyl methacrylate, or 1,4-butanediol, dimethacrylate, poly(ethylene glycol) dimethacrylate, poly(ethylene glycol) diacrylate, polyethylene oxide dimethacrylate, polyethylene oxide diacrylate, N,N'-methylenebisacrylamide, N,N' The crosslinking agent may be, for example, methylene bis(2-methylacrylamide), methylene dimethacrylate, N,N'-methylene bis(2-methylacrylamide), methylene diacrylate, methylene bis(2-methylacrylate), diethylene glycol diacrylate, hexamethylene diacrylate, oxybis(methylene)bis(2-methylacrylate), and oxybis(ethane-2,1-diyl)bis(2-methylacrylate), as well as crosslinkers based on metal salts (e.g., sodium chloride, calcium chloride, calcium nitrate, etc.). In certain embodiments, the hydrogel precursor solution includes acrylic acid and N,N'-methylene bisacrylamide. The solvent used in the hydrogel precursor solution may include water, alcohol, and combinations thereof.

[0067] The hydrogel precursor solution can be deposited onto one or more patterned metal features using methods known in the art for depositing solutions onto substrate surfaces. Exemplary deposition methods include, but are not limited to, spin coating, printing, print screening, spraying, painting, doctor blading, slot-die coating, or dip coating.

[0068] Methods for curing the hydrogel precursor coating can vary based on the selection of the hydrophilic polymer and, optionally, the crosslinker, hi certain embodiments, the curing method comprises heating, irradiation with electromagnetic radiation (e.g., ultraviolet light), reaction with a radical initiator (e.g., azobisisobutyronitrile, etc.), or a combination thereof.

[0069] The patterned mask can be a reflective mask (such as those used in extreme ultraviolet (EUV) lithography), a transmissive mask, a binary intensity mask, a phase shift mask, a quartz mask, as well as other mask types known in the art.

[0070] The developer removes the exposed portions of the resist layer (e.g., the exposed portions of a positive photoresist) or the unexposed portions of the resist layer of a negative photoresist, thereby forming a patterned photoresist layer. Examples of suitable developers for negative photoresists include aqueous alkaline solutions (e.g., diluted sodium hydroxide, diluted potassium hydroxide, etc.) or aqueous solutions of metal-ion-free organic tetramethylammonium hydroxide (TMAH).

[0071] The method of depositing the metal is not limited to any particular method. Accordingly, all known methods of deposition are contemplated by this disclosure. In certain embodiments, the metal is deposited by physical vapor deposition, chemical vapor deposition, sputtering, atomic layer deposition, electrochemical deposition, spin coating, dip coating, inkjet printing, and any combination thereof.

[0072] Removing the irradiated photoresist layer can include contacting the irradiated photoresist layer with a stripping solvent. The stripping solvent can be any solvent in which the irradiated photoresist layer is at least partially soluble. Exemplary stripping solvents include, but are not limited to, organic solvents (e.g., dimethyl sulfoxide, acetone, N-methyl-2-pyrrolidone, etc., and mixtures thereof).

[0073] In cases where a flexible substrate is plasma-treated, the time for which the flexible substrate is plasma-treated can vary based on the characteristics of the biocompatible elastomeric fiber and the plasma treatment conditions. Selecting appropriate plasma treatment conditions is well within the capabilities of one skilled in the art. In certain embodiments, the plasma treatment of the flexible substrate is performed for 0.1 to 60, 0.5 to 60, 0.5 to 55, 0.5 to 50, 0.5 to 45, 0.5 to 40, 0.5 to 35, 0.5 to 30, 1 to 30, 5 to 30, 10 to 30, 15 to 30, 20 to 30, 25 to 30, 5 to 25, 5 to 20, 5 to 15, 5 to 30, 5 to 10, 5 to 25, or 10 to 20 minutes.

[0074] The sacrificial layer can be removed by contacting it with an aqueous solvent, thereby at least partially dissolving the sacrificial layer. In certain embodiments, at least 90%, at least 95%, at least 96%, at least 97%, at least 98%, at least 99%, at least 99.9%, or all of the sacrificial layer is removed by dissolution in the aqueous solvent.

[0075] The method for separating the substrate and the flexible substrate including one or more patterned metal structures disposed on the surface of the flexible substrate is not particularly limited. Therefore, any separation method known in the art is contemplated by the present disclosure. In a specific embodiment, the substrate and the flexible substrate including one or more patterned metal structures disposed on the surface of the flexible substrate are physically separated.

[0076] The liquid metal can be deposited onto one or more patterned metal structures by selective wetting, converting the one or more patterned metal structures into one or more electronic components, thereby forming a permeable stretchable bioelectronic device.

[0077] Biocompatible elastomeric fiber flexible substrate Fabrication of transparent and stretchable liquid metal microelectrodes (μLMEs) Figure 1a shows a schematic illustration of the wafer-scale μLME fabrication process, which consists of four major steps: 1) photolithography of Ag on a SiO2 wafer pre-modified with a thin layer of water-soluble dextran; 2) electrospinning of a fibrous poly(styrene-block-butadiene-block-styrene) (SBS) mat onto the Ag micropattern; 3) dissolution of the dextran layer and transfer of the Ag micropattern from the SiO2 wafer to the SBS fibrous mat, and 4) selective wetting of LM on the Ag-coated areas and generation of μLME.

[0078] Importantly, direct electrospinning of SBS fiber mats onto Ag micropatterns is a key step that addresses two crucial challenges in fabricating microscale LMs on porous, stretchable substrates (Figure 1b). First, direct electrospinning of SBS fibers ensures conformal contact between the flexible substrate and the Ag micropattern. Therefore, when the dextran layer dissolves in water, the Ag micropattern adheres tightly to the SBS fiber mat, achieving complete pattern transfer. Second, the fibrous structure of the host SBS substrate provides a large interface for rapid water penetration and the water / dextran contact interface, thereby achieving fast and isotropic dissolution of dextran. Furthermore, such isotropic dissolution of the sacrificial layer is crucial for achieving complete pattern transfer over a large area (Figure 6).

[0079] We choose eutectic gallium-indium (EGaIn) alloy as the LM. EGaIn exhibits a low melting temperature of 15.4 °C and stable conductive properties even under large tensile strains due to its inherent fluidity compared to solid electrodes. In addition, the low elastic modulus and high biocompatibility of EGaIn also benefit applications in soft and stretchable bioelectronics.

[0080] As a proof-of-concept, we patterned 300-nm-thick Ag with feature sizes ranging from 2 μm to 1,000 μm on a 4-inch wafer substrate (Figure 2a, inset). The SiO2 wafer surface was pre-spin-coated with a 10-nm-thick, smooth dextran layer (Figure 7). Next, a 100-μm-thick SBS fiber mat was electrospun onto the wafer surface (Figure 8). After plasma treatment of the SBS side, the entire sample was immersed in a water bath. Water rapidly wetted the sample through the micropores, dissolving the dextran layer, and a free-standing Ag-coated SBS fiber mat was obtained within minutes (Figure 2a). After drying, EGaIn was applied onto the SBS substrate. Due to its poor affinity for SBS and high reactivity with Ag, EGaIn selectively dewetted the SBS surface while wetting the Ag-covered areas, forming a μLME on the ultrasoft SBS fiber mat (Figure 2b).

[0081] Importantly, the success rate of pattern transfer largely depends on the water penetration rate through the SBS fiber mat. After electrospinning, the SBS fiber mat was initially hydrophobic, exhibiting a large water contact angle of 130° (Figure 9). At this point, it took 2 min to wet the bottom of the substrate, and the dextran layer was difficult to dissolve. After plasma treatment (Figure 2c), the SBS surface became superhydrophilic. With a plasma time of 20 min, the penetration time decreased to 3 s (Figure 10). Therefore, the pattern transfer rate (defined as the area with transferred Ag features relative to the total patterned area) increased from nearly zero (before plasma treatment) to over 95% (Figure 2d, Figure 11). In contrast, samples coated with impermeable spin-coated SBS films showed no trace of pattern transfer even after immersion in water for 1 week (Figure 12).

[0082] Note that the selective wetting of EGaIn relies on the reactive alloying between Ag and In to form an AgIn alloy (which exhibits a high affinity for additional EGaIn applied on top of the Ag area) (Figure 13). As the Ag layer increased from 0 to 300 nm, the EGaIn contact angle decreased from 140° to 30° (Figure 2e). We obtained μLMEs with dimensions ranging from 2 μm to 1,000 μm and a wide variety of shapes, including arbitrary patterns corresponding to various electrical components (lines, dots, source / drain electrodes, interconnects, mesh electrodes, antennas, generators, inductors, capacitors, and resistors) (Figures 14 and 15). Compared to the original Ag patterns, slight enhancement of the EGaIn characteristics was observed. For example, as shown in Figure 2f, the line width of the Ag electrode on the SBS fiber mat was 3.2 μm. After wetting, the final EGaIn electrode was 3.3 μm. Note that the pattern pitch remained unchanged. The linewidth of the μLME showed a linear relationship with that of the Ag, with a slope of approximately 0.99 (Figure 2g). Figure 2h shows an electron patch based on the μLME, consisting of an arbitrary polygon, and Figure 2i shows an electron patch with an electrode density of 75,500 electrodes / cm. 2 This shows high density electronic components.

[0083] Permeability, biocompatibility, conductivity, and stretchability of μLME Due to the high porosity of the fiber mat, μLME has excellent permeability to air, moisture, and liquid. Figure 3a shows the gradual penetration of a colored liquid from the top side to the bottom side of the μLME, demonstrating its excellent liquid permeability. The air permeability of μLME on a 330 μm thick SBS fiber mat is 78 mm / s, and the moisture permeability is 835 g / m. 2 / day (Fig. 3b, Fig. 3c). In contrast, the air permeability of Ecoflex and PDMS thin films was almost zero. Their water permeability was 50 g / m 2 / day. Furthermore, the permeability of μLME can be further enhanced by decreasing the thickness of the SBS. At a thickness of 25 μm, the air permeability was 235 mm / s and the moisture permeability was 990 g / m 2 / day, respectively. In addition, in an in vitro study using L-929 cells as a model cell, μLME demonstrated low cytotoxicity (Figure 16). Furthermore, in an in vivo animal study conducted on rabbit skin, no obvious inflammation was observed (i.e., no erythema or edema on the skin) during the observation period (24, 48, and 72 hours) (Figure 17).

[0084] We characterized the conductivity of μLME at different line widths. When the line width was increased from 2 μm to 200 μm, the μLME was 1.3–3.9 × 10 5 This was associated with high levels of conductivity in the range of S / cm (Fig. 3d, Fig. 18). We note that the conductivity of the μLME is higher than that of EGaIn alone (but lower than that of Ag alone). This enhancement can be attributed to the contribution of a more conductive Ag layer underneath the LM. It is also noted that in the absence of LM, the transferred Ag on the SBS fiber mat cannot withstand tensile strains greater than 50% due to the presence of cracks ranging from a few micrometers to tens of micrometers (Fig. 19).

[0085] In contrast, μLMEs were highly stretchable and conductive. The resistance of 5 μm μLMEs increased by only 0.6% at 300% tensile strain (Figure 3e). When the line width was increased to 10 and 50 μm, μLMEs exhibited electrical stability at higher tensile strains, with the resistance increasing by only 1.3 and 1.1 times, respectively, at 1,000% strain. When the line width was increased to 200 μm, the resistance changed by only 5.25 times at 1,500% strain. Furthermore, μLMEs with various line widths (10, 50, and 200 μm) exhibited electrical stability and robustness in stretch-relaxation cycle tests under large strains (over 1,000% strain) due to the formation of reversible buckling structures (Figure 21) (Figure 3f, Figure 20). The resistance of the μLME changed by only 1.03 times (10 μm line width) after 100 cycles of stretch-relaxation testing at 1,000% strain, and by only 0.393 times (50 μm line width) after 500 cycles.

[0086] μLME-based implantable bioelectronics for neural interfaces with long-term biocompatibility Due to its properties of ultrasoftness, biocompatibility, stretchability, and high spatiotemporal resolution, μLME is suitable for implantable bioelectronics, where high device density and long-term comfort are crucial. As a proof-of-concept, we fabricated implantable μLME arrays to record ECoG signals arising from soft, curved, and delicate neural interfaces. Major subdomains of the rat cerebral cortex (composed of the motor cortex, somatosensory cortex, visual cortex, and retrosplenial cortex) were coated with μLME arrays (Figure 4a). The μLME arrays feature small-diameter circular electrode units (500 μm), fine interconnects (40 μm), and a density of 100 electrodes / cm. 2 The LMEs were characterized by a high channel density of 100 μm (Fig. 22). Based on the similarity in mechanical compatibility with brain tissue (Fig. 23), the μLMEs (25 μm thick) conformally adhered to the cortical surface (Figs. 4b and 4c), which exhibited a much lower Young's modulus (Fig. 4d) and superior conformality (Fig. 24) compared with reported electrode materials for ECoG devices (e.g., Au / PI, Au / Parylene, and silicon).

[0087] To verify the potential of the μLME array for neural electrophysiological recording and stimulation, we first performed in vitro electrochemical impedance characterization. The electrochemical impedance of the μLME interface in phosphate-buffered saline (PBS) was similar to that of commonly used Au / PI electrodes with identical dimensions in the frequency range of 1 Hz to 10,000 Hz. The μLME array was easily stretchable up to 500% and maintained good stability in impedance and phase signals, whereas the Au / PI electrodes showed dramatic changes when stretched up to 30% (Figures 4e and 25). Furthermore, the 36-channel μLME array showed small variations in electrochemical impedance between individual electrodes (Figure 26).

[0088] To record cortical activity in vivo, we recorded neural signals from sleeping rats (Figure 27). Signal power spectra revealed non-rapid eye movement (NREM) sleep, associated with low-frequency delta waves below 4 Hz and theta waves in the 4-8 Hz range (Figure 4f). To verify the region-specific recording capabilities of the μLME, we stimulated the forelimb with brief current pulses and examined the somatosensory evoked potentials (SEPs) elicited within the cortex. SEPs were recorded during electrical stimulation at frequencies ranging from 1 Hz to 9 Hz (Figure 4g) and pulse voltages ranging from 1 V to 6 V (Figure 4h). The amplitude of the neural signals exhibited intensity-dependent behavior, ranging from 15 to 107 mV for positive SEPs and 44 to 630 mV for negative SEPs with increasing pulse voltage. Notably, because these electrical stimuli were applied to the right forelimb, neural signals in the contralateral (left) hemisphere were much stronger than those in the ipsilateral (right) hemisphere (Fig. 4h). Furthermore, a rhythmic EEG signal (7–14 Hz) was reproducibly observed throughout 10 consecutive electrical stimuli (Fig. 4i).

[0089] Figure 4j shows the spatiotemporal characteristics of the μLME array. In response to forelimb stimulation, positive SEP peaks (red region) were detected in the stimulated hemisphere approximately 20 ms after stimulation onset in both forelimbs, while these responses were absent in the contralateral hemisphere. In contrast, negative SEP peaks (blue region) appeared in both hemispheres, indicating a cognitive response to stimulation. Other cortical regions (e.g., motor (primary and secondary) cortices and visual cortex) also showed weaker SEP signals compared to the somatosensory cortex.

[0090] The μLME array demonstrated excellent long-term biocompatibility, which is important for long-term implantation. Immunohistological analysis of brain slices revealed no tissue or cellular damage or inflammatory response immediately after implantation (Figure 5a). Two weeks after implantation, microglial activation levels in the μLME group showed no significant differences compared to the control group in both cell body size and Iba1 count (Figure 5b). In contrast, implantations using Au / PI and Au / PDMS for two weeks showed significant increases in microglial cell body size and Iba1 abundance compared to the control (p<0.001, t-test), indicating significant microglial activation and inflammatory response (Figures 5c-d). Finally, compared to current ECoG devices, the implanted μLME exhibited long-term biocompatibility in living rats for over four months (Figure 5c). μLME also demonstrated superior biocompatibility to Au / PI and Au / PDMS when these samples were implanted into rat skeletal muscle for 1 week (Figure 28). This advantage is most likely due to the improved softness, stretchability, and permeability of μLME.

[0091] Porous and stretchable substrates (e.g., electrospun fiber mats) have excellent transparency, flexibility, and biocompatibility—important properties for bioelectronic applications. However, unlike currently mainstream planar substrates (e.g., thin-film plastics or elastomers) (Figure 36), patterning high-resolution electrodes on rough and porous substrates has been a challenging task. The patterning process is primarily carried out by printing methods (e.g., screen printing or inkjet printing) or shadow mask techniques. Therefore, patterning resolution has been limited to 50–10,000 μm, and device density has been very low (Figure 37).

[0092] The successful preparation of μLMEs on large, transparent, and stretchable fiber mats will pave the way for high-density, integrated, implantable, and long-term biocompatible LM electronics. Currently, μLMEs have been fabricated by laboratory-based photolithography with a typical patterning resolution of 2 μm and a density of 990 g / m. 2 / day moisture permeability, 75,000 electrodes / cm 2 Device density exceeding 10 5 They exhibit conductivity on the order of S / cm and stretchability up to 1,500%. In principle, the patterning resolution of μLME can be further enhanced to the submicrometer (0.25-1 μm) range by using, for example, high-resolution lithography tools and fiber mats consisting of smaller diameter fibers (Figure 29).

[0093] Hydrogel flexible substrate The μLMH allows for tight ion-electron interactions with free ion diffusion between the LM and the hydrogel. At the biointerface between electronics and biology, electron-ion exchange occurs in the electrode / extracellular space, enabling bidirectional communication (recording and stimulation). Compared with either wet gel electrodes or dry contact electrodes, the μLMH allows for tight ion-electron interactions with high electrical conductivity, enabling bidirectional communication with low impedance and high signal fidelity (Figures 39b and 45). Spreading resistance (Rs) is the resistance of the interconnects leading to higher-level circuits (e.g., amplifiers). On the other hand, in the case of μLMH, Rs is negligible due to its metallic properties, making it an excellent candidate for low-impedance biointerfaces. A low-density μLMH with a resolution of 5 μm enabled a transparent and seamless biointerface integrated with curved, soft, stretchable human skin (Figure 39d). The μLMH also achieved a density of 2500 electrodes / cm. 2 For stretchable circuits, it can be patterned into a variety of high-density shapes and dimensions, including lines, dots, wires, and interconnects (Figure 39e, Supplementary Figure 46).

[0094] Low impedance μLMH due to tight ion-electron interactions The μLMH has low and stable impedance due to its tight ion-electron interaction and the excellent properties of the LM (e.g., high elasticity and electrical conductivity). The tight ion-electron interaction of the μLMH was confirmed by time-of-flight secondary ion mass spectroscopy (TOF-SIMS) 3D profile of the ion distribution of Ga and In ions in the μLMH (Figure 40a). This ion-electron interaction is due to the metal ions (mainly Ga) being ionized, as evident in the X-ray photoelectron spectroscopy (XPS) spectra (Figures 40b-40d) and the red shift of the Fourier transform infrared spectroscopy (FTIR) spectrum of the PAA / LM (Figure 40e). 3+ and In 3+ This was attributed to the coordination between the carboxyl groups of PAA and the hydroxyl groups of hydroxyapatite.

[0095] Compared with inert metals (e.g., Au), LM, acting as a contact electrode, provides an interfacial impedance much closer to that of the hydrogel alone due to its larger charge injection capacity (Figure 40f). LM interacted differently with hydrogels with various ionic conductivities. In the case of LM / PAA / LM (ionic conductivity of approximately 0.66 S / m), a soft circuit occurred, indicating an electrochemical corrosion process between LM and PAA. In the case of LM / PAA2 / LM (ionic conductivity of approximately 2.73 S / m), the Nyquist plot showed that the electrode impedance was modeled as a general Warburg impedance, suggesting the existence of both charge transfer and ion diffusion processes at the LM / PAA interface (Figure 40g). The various ionic conductivities of PAA were driven by LM, which exhibited negligible signal distortion (Figure 2h, Figure 47). The μLMH exhibited much lower skin impedance than commercial gel electrodes (Ag / Ag / Cl), metal electrodes (LM / PDMS), and wet electrodes with LM and a conventional metal (Au) as the current collector (Figure 40i). Furthermore, the electrochemical impedance of the μLMH showed superior performance to other biological interfaces, including dry metal (Au) electrodes, LM electrodes, and hydrogel-only electrodes, in terms of impedance, maximum distortion, and electrochemical stability (Figure 2j, Figure 48).

[0096] High-resolution patterning of μLMHs resulting from tuning interfacial adhesion Here, we tuned the adhesion between the two interfaces during the gelation process. Generally, the transfer printing process consists of collecting / picking up the metal ink from the donor substrate and printing / delivering the ink to the receiving substrate (steps 1 and 3 in Figure 49). First, with increasing gelation time, the hydrogel stamp came into stronger contact with the aligned, peelable micropattern ink on the Si2 wafer (donor substrate) through bonding between hydrophilic groups (e.g., hydroxyl and carboxyl groups). Meanwhile, the water-soluble sacrificial layer was gradually dissolved by the hydrogel precursor solution. Furthermore, the as-formed hydrogel matrix is ​​water-permeable; therefore, even when the top of the micropattern is covered by the hydrogel, the sacrificial layer can be continuously dissolved by water. The dissolution rate of the sacrificial layer was approximately 1 nm / s for a 10 nm-thick coating (Figure 50). Therefore, during the gelation period, the adhesion at the stamp / ink interface was stronger than that at the ink / substrate interface during the collection period, but weaker during printing. After peeling the hydrogel from the wafer, the micropatterns were selectively wetted with LM (eutectic gallium-indium (EGaIn) alloy) to generate μLMH.

[0097] As a proof-of-concept, polyacrylic acid (PAA) hydrogel was employed due to its hydrophilic surface binding, high adhesiveness, good biocompatibility, and mechanical properties. 30The preparation of the hydrogel-receiving substrate is shown in Figure 51. To maintain a relatively stable water content, an organic solvent (glycerin) was introduced to form an organohydrogel. Here, the crosslinking density of PAA increased with increasing crosslinking duration, and a sol-gel transition was observed at approximately 58 s (Figure 41a). Meanwhile, the interfacial adhesion strength increased with increasing gelation duration, triggering pattern transfer (Figure 41b). AA was fully crosslinked and polymerized to PAA after approximately 60 s of UV exposure, as confirmed by the chemical shift in proton nuclear magnetic resonance (NMR) spectroscopy and the enhanced C=O stretching signal and the attenuated C=C stretching signal in Raman spectra (Supplementary Figure 51). At this point, the pattern transfer rate reached approximately 96% (Figure 3c, Supplementary Figure 52). The as-prepared PAA film exhibited high shear adhesion strength when attached to various substrates and organs (Supplementary Figure 47). After the transfer of the Ag micropattern to the organohydrogel and selective wetting with LM, transparent LM micropatterns were obtained in the form of a mesh structure with a typical resolution of about 9 μm (Figure 41d). The transferred LM microelectrodes show no obvious difference in resolution. With increasing storage time, the line width can be further narrowed to 8.9 μm due to the moderate dehydration of the organohydrogel substrate with a certain mass fraction of glycol in the precursor solution (Figure 54). At this line width, the mass loading of LM during the selective wetting process ranged from about 0.91 to 4.09 mg / cm. 2 (Figures 41e, 41f, and 55).

[0098] Stretchability, transparency, and electrical conductivity of μLMH The μLMH exhibits excellent stretchability and flexibility. Increasing the mass fraction of the precursor solution (water / glycerin containing AA) to 30 wt% decreased the μLMH's elastic modulus to approximately 25 kPa (Figure 42a), and the maximum strain reached approximately 2,500% (Figure 42b). Furthermore, the transmittance of the μLMH increased with decreasing patterning linewidth, with an average transparency of approximately 89.8% at 550 nm (Figure 42c). Furthermore, the electrical conductivity increased with increasing LM mass loading, with a maximum conductivity of approximately 31,967 S / cm (Figure 42d). During the stretch-relaxation test, the μLMH exhibited good electrical stability. For a highly conductive μLMH (50 μm linewidth and gap, average transmittance of 53.7 at 550 nm), the electrical resistance only changed by approximately 3.8 with 1,000% strain (Figure 42e). For a highly transparent μLMH (10 μm line width and gap, average transmittance of 79.8 at 550 nm), the electrical resistance changed only about 2.1 with 1000% strain (Figure 56). The μLMH can also function as a stretchable analog circuit integrated with an LED and an amplifier for power management (Figure 42f).

[0099] μLMH for transcutaneous electrical stimulation and electrophysiological sensing with high SNR and fast response A low-impedance μLMH offers advantages for biointerfaces, combining both a high SNR and a short response time. Such low-impedance μLMHs are crucial for highly sensitive and fast-response biointerfaces. When various biointerfaces were applied for transcutaneous electrical stimulation at various intensities (Figure 43a) and frequencies (Figure 43b), the μLMH demonstrated significantly higher charge injection efficiency into the skin, resulting in higher current intensities compared to other biointerfaces. Additionally, the corresponding electromyogram (EMG) signals recorded by the μLMH from the adjacent forearm region had higher signal quality in terms of a higher SNR (Figure 43c) and shorter response time (Figure 43d) under identical electrical stimulation conditions (1 Hz, 20 V, Figure 43e). Next, EMG signals were recorded using the μLMH under various electrical stimulation frequencies ranging from 1 Hz to 100 Hz (Figure 43f). The spectrogram of the EMG signal showed a continuous, clear, and complete response to fast electrical stimulation (100 Hz) (Figure 43g).

[0100] μLMH for neural interfaces with high spatiotemporal optogenetic mapping capabilities Due to this wafer-based patterning technique and its mechanical compatibility with brain tissue (25 kPa), μLMHs were successfully prepared on hydrogel surfaces with much higher resolution than those achieved using conventional techniques (Figure 60). We demonstrated an electrocorticography (ECoG) array with optogenetic mapping capabilities for recording and intervening in flexible, curved, and precise neuroelectrophysiological interfaces (Figures 44a and 44b). We developed a multilayered ECoG device consisting of a PAA substrate, 32-channel μLMH electrodes, wiring, and a polyacrylamide (PAM) hydrogel packaging layer to prevent signal crosstalk (Figure 44c). Major cortical subdomains in rats, consisting of the motor cortex, somatosensory cortex, visual cortex, and retrosplenial cortex, were coated with μLMHs (Figure 57). The circular electrode array had a diameter of 600 μm, a minimum line width of 200 μm, and 32 channels.

[0101] To validate the potential of the μLMH for neuroelectrophysiological recording and stimulation, we first performed in vitro impedance characterization. In multichannel testing, the 32-channel μLMH electrode array exhibited low electrochemical impedance along with low device variability (Figure 44d).

[0102] To record cortical activity in vivo, we recorded neural signals from sleeping rats (Figure 58). To verify the ability of the μLMH to record region-specific cortical activity using optogenetic mapping, we stimulated the cortex using 473 nm blue light and examined somatosensory evoked potentials (SEPs) in the cortex. Notably, neural signals with light stimulation were twice as strong as those without light stimulation (Figure 59). Furthermore, the rhythm of the cortical EEG signal (0–50 Hz) was reproducibly observed throughout continuous light stimulation (Figure 6e). The spatiotemporal characteristics of the SEPs were also investigated, and the SEP peak was observed approximately 80 ms after stimulation onset (Figure 44f). The μLMH demonstrated excellent biocompatibility, with no significant difference between the μLMH group and the control group (no implant, Figures 44g, 44h).

[0103] Bioelectronics that interface with the human body are fundamental to the rapidly developing fields of physiological monitoring, imaging, neuromodulation, neuroscience, and tissue repair. The ability of flexible iontronics to achieve both electronic and ionic conduction makes them excellent candidates for biointerfaces. However, current flexible iontronics rarely achieve high electrical and ionic conductivities, low electrode / electrolyte interfacial impedance, and high integration density.

[0104] Here, we developed wafer-scale patterned LM microiontronics on hydrogels (μLMHs) capable of both electronic and ionic conduction with patterning resolution up to 5 μm. μLMHs exhibit tissue-like softness (approximately 25 kPa), a maximum strain of 2,500%, and low and stable interfacial impedance under stretching. To highlight its versatility, we use this platform to demonstrate transcutaneous electrical stimulation and electrophysiological sensing capabilities with both a high signal-to-noise ratio (SNR) and a short response time. Furthermore, we demonstrate a multichannel neural interface with high-resolution spatiotemporal optogenetic mapping.

[0105] method Materials. Solvents, including acetone, tetrahydrofuran, dimethylformamide, ethanol, and 2-propanol, were purchased from Anaqua Global International Inc. Limited. Negative photoresist (NR9-1500P) and developer for NR9-1500P (DR6) were obtained from Futurrex, Inc. (USA). Liquid metal and dextran were purchased from Sigma-Aldrich and used as received. All materials were used as received.

[0106] Biocompatible elastomeric fiber flexible substrate Fabrication of transparent and stretchable liquid metal microelectrodes (μLMEs). Si or SiO2 wafers were first sonicated in acetone, 2-propanol, and deionized (DI) water, respectively, and then blown dry with nitrogen gas. A sacrificial layer was then formed by spin-coating a dextran solution (10 wt % aqueous solution) onto the wafer at 4,000 rpm for 40 seconds, followed by baking on a hotplate at 80 °C for 1 minute and 180 °C for 30 minutes. The baked dextran-coated wafers were then spin-coated with a layer of negative photoresist NR9-1500P (24%–28% solids, primarily cyclohexanone) at 4,000 rpm for 40 seconds, followed by pre-baking on a hotplate at 155 °C for 1 minute. The photoresist was then applied at 170 mJ / cm2 with the aid of a mask aligner (Suss MA6). 2 The wafer was exposed to UV light at a dose of 1000 s. After UV exposure, the photoresist was post-baked at 105 °C for 3 minutes and then developed with DR6 developer for 10–15 seconds. After rinsing the sample in DI water and drying with compressed nitrogen gas, a 300 nm thick Ag layer was deposited on the wafer by thermal evaporation. A peeling process was then carried out in acetone solvent to form Ag micropatterns on the wafer. To transfer the Ag micropattern to a stretchable, transparent, flexible substrate, a polystyrene-block-butadiene-block-styrene (SBS) fiber mat with a thickness ranging from 25 to 330 μm was electrospun directly onto the wafer surface with the Ag micropattern deposited, followed by air plasma treatment (Harrick) for 20 minutes. Next, the sample was immersed in DI water at room temperature for up to several minutes to dissolve the sacrificial layer between the wafer and the Ag micropattern. The SBS fiber mat was peeled off from the wafer, and the Ag micropattern was transferred to the fiber mat. To obtain the μLME, LM was applied to the Ag micropatterns by selective wetting in an Ar gas-filled glove box.

[0107] Measurement of the water penetration rate through an SBS fiber mat. The water penetration rate through an SBS fiber mat is a reflection of the water contact angle, wetting time, wetting radius, and absorption rate. The contact angle of water on the plasma-treated SBS fiber mat for different periods was measured by a contact angle meter (SDC-350). The amount of water droplet was fixed at 5 μL. The wetting time, wetting radius, and absorption rate of water were measured by a moisture management system (MMT, SDL Co.) according to the standard AATCC79.

[0108] Characterization of the morphology and permeability of μLME. The morphology and surface roughness of the sacrificial layer were characterized using an atomic force microscope (AFM, XE70). The line width and thickness of the μLME were determined using a 3D optical surface profiler (New Zygo NexView). The surface morphology and elemental mapping of the samples were collected using a scanning electron microscope (SEM, TESCAN VEGA3). Air permeability tests were performed according to the ASTM D737-08 standard using an MO21S air permeability tester (SDL Americ, Inc.). The air flow pressure was set to 100 kPa. Moisture permeability tests were performed using the cup method according to standard E96 / E96M-13. The test duration was 72 hours. Both air permeability and moisture permeability tests were performed at a constant temperature (22 °C) and humidity (63%).

[0109] Electrical and Mechanical Characterization of μLMEs. The electrical conductivity (σ) of μLMEs was calculated according to the formula σ = L / (A × R), where L, A, and R are the length, cross-sectional area, and electrical resistance of the μLME, respectively. The A of μLMEs was determined using a surface profiler (New Zygo NexView) (Figure 14), and the integrated polygon area was calculated using the software OriginPro 8.5. The R of μLMEs (L = 10 mm) with different line widths was measured by characterizing the direct current (DC) IV using a parameter analyzer (Keithley 4200A-SCS parameter analyzer) connected to a probe station (Micromanipulator) (Figure 18). The electrical resistance of μLMEs under different strains was measured by the four-probe method using a source meter (Keithley 2400) connected to a customized stretcher (Precise).

[0110] Electrochemical Characterization. Electrochemical impedance spectroscopy (EIS) was performed using an electrochemical workstation (CHI 660E) equipped with a three-electrode testing system. Ag / AgCl (2 M KCl), platinum foil, and μLME were used as the reference, counter, and working electrodes, respectively. 1× phosphate-buffered saline (PBS) solution (Gibco) was used as the electrolyte. Impedance and phase angle were recorded as a function of frequency in the range of 1 Hz to 10,000 Hz with an AC sinusoidal signal amplitude of 10 mV. EIS spectra of μLME stretched at various strain rates were acquired using the same electrochemical workstation coupled with a homemade stretching device.

[0111] Rat Surgery. All animal experiments and procedures following the experimental protocol were reviewed and approved by the Animal Research Ethics Sub-Committee, Research Committee at the City University of Hong Kong (approval number A-0664). Male Sprague-Dawley rats (7-8 weeks old, 200-300 g) were ordered from the Laboratory Animal Services Centre at the Chinese University of Hong Kong. All rats were housed under a 12:12 h light / dark cycle in the Laboratory Animal Research Unit at the City University of Hong Kong until the experiment. Before surgery, rats were first anesthetized with gaseous isoflurane (air mixed with 3% isoflurane, 2 L / min) and then deeply anesthetized with sodium pentobarbital (30 mg / kg). After anesthesia, the rats' heads were shaved at 37°C. First, a midline incision was made above the craniovertebral junction. The overlying muscles were then cut and retracted to expose the skull. A craniotomy was performed using a dental drill, and the dura was then peeled away using fine forceps to ensure complete exposure of the cerebral cortex.

[0112] Electrocorticography (ECoG) neural signal recording and electrical stimulation. The μLME array, coupled to a flexible printed circuit board (FPC) connector, was connected to the PCB using a customized FPC adapter. ECoG neural signals were then recorded using a multichannel high-precision data acquisition system (PowerLab 16-35, AD Instruments). Signals were acquired and digitally bandpass filtered (1–30 Hz) using the corresponding software, LabChart. Continuous electrical stimulation was performed on both the left and right forelimbs of the rats using square-wave voltage pulses with various applied frequencies (1, 3, and 9 Hz) and intensities (1, 3, and 6 V) generated from the stimulation output port of the PowerLab through a BNC cable. A total of 50 pulses for each stimulation condition were delivered at a frequency of 1 Hz, and 50 responses per stimulation condition were averaged. Custom-written Matlab code was used to analyze the ECoG neural signal spectrograms.

[0113] Cryosectioning, histological staining, and observation of brain slices. After implantation of the μLME array, the histological condition of the brain slices was first examined. To preserve the tissue structure, the samples were fixed in PBS containing 4 wt% formaldehyde for 24 hours, transferred to a 30% sucrose solution overnight, and then transferred to a sucrose-Cryo-OCT compound before cryosectioning. The brain samples were then frozen at -80°C and sliced ​​into 15-μm-thick blocks using a cryomicrotome (CryoStar NX70). Hematoxylin and eosin (H&E) staining was performed using the basic dye hematoxylin (BL735A-1) to stain acidic cellular components (i.e., nucleic acids, glycosaminoglycans, and acidic glycoproteins) and the acidic dye eosin (BL735B) to stain the cytoplasm of cells.

[0114] Specifically, after washing with DI water, the slices were first stained with hematoxylin solution for 5 minutes. Next, the slices were thoroughly rinsed with DI water, differentiated in an aqueous solution containing 0.1% acetic acid and 85% ethanol for 10 seconds, and then rinsed again with DI water for 1 minute. Next, the slices were stained blue in a PBS / PBST solution for 30 seconds and then rinsed with DI water. After the bluing, the slices were washed in ethanol for 10 seconds, and eosin was applied to further stain the tissue for 30 seconds. Finally, the slices were dehydrated in ethanol (95%) and permeabilized twice in xylene (5 minutes each). For immunohistochemical analysis, the slices were observed under an inverted microscope (Nikon Eclipse Ti).

[0115] Cytotoxicity Test of μLME. Samples for cytotoxicity testing included a control sample (fresh complete medium), absorbent gauze, SBS mat, μLME, and 20% dimethyl sulfoxide (DMSO). L-929 cells (ATCC, USA) were used to investigate the in vitro cytotoxicity of all samples according to the ISO 10993-5: 2009(E) standard. All L-929 cells were cultured at 6 × 10 cells in complete Dulbecco's Modified Eagle's Medium (DMEM, Hyclone, Hong Kong) supplemented with 10% fetal bovine serum (FBS, Hyclone, Hong Kong) and 1% penicillin / streptomycin (PS, Hyclone, Hong Kong) in a 24-well plate. 4 The cells were seeded at a seeding density of 100 cells / mL. After 24 hours of incubation and confirmation of cell confluence and morphology, the medium was discarded and 1 mL of fresh complete medium was added to the culture plate. Next, 0.2 cm 2 Sterile samples were added to 24-well plates to evaluate the cytotoxicity of the materials. All samples were incubated at 37°C in an incubator containing 5% CO2.

[0116] After discarding the culture medium, 200 μL of a staining solution containing PBS, 2 μM calcein AM, and 8 μM propidium iodide was added and incubated for 45 minutes. The cells were then rinsed with PBS and observed under a fluorescence microscope (Zeiss, Germany). To examine cell proliferation, 500 μL of MTT solution was added to each well and incubated for 2 hours. Next, 500 μL of DMSO was added to dissolve the formazan. Absorbance at 450 nm, representing the metabolic activity of the cells, was measured using a plate reader (BioTek, USA). Cell viability was calculated using the absorbance ratio between the test wells and the blank control wells.

[0117] In vivo evaluation of irritation and skin sensitization potential of μLME. Animal testing of samples was performed according to the ISO 10993-10:2010(E) standard. All procedures followed the ethical guidelines for laboratory animals approved by The Hong Kong Polytechnic University. Four groups of samples were evaluated, including a negative control (thin cotton cloth) as group 1, a positive control (thin cotton cloth rinsed with saturated sodium dodecyl sulfate) as group 2, an SBS mat as group 3, and a μLME with a serpentine structure as group 4. All samples were measured on 2.5 × 2.5 cm plates. 2 The samples were then attached to New Zealand white rabbits (2.0-3.0 kg body weight). The rabbits' backs were shaved 24 hours before the sample attachment. Each group had three sample patches, which were attached to a gauze pad (2.5 × 2.5 cm) secured with a semi-occlusive medical sterile wound dressing (Hynaut Co., Ltd., Qingdao, China). 2 ) was applied to the skin beneath the skin. After 24 hours of exposure to shaved skin, the patch was removed. The test site was then observed for visible changes (e.g., erythema) at 0 (initial), 24, 48, and 72 hours after sample application. The mean erythema score (MES) was measured on a scale of 0 to 4, with grade 0 representing no erythema, grade 1 representing slight erythema, grade 2 representing moderate erythema, grade 3 representing moderate to severe erythema, and grade 4 representing severe erythema.

[0118] Long-term biocompatibility testing of implanted μLME. For biocompatibility testing in the brain, electrodes (μLME, Au / PI, and Au / PDMS with identical electrode designs) were implanted into the cerebral cortex to evaluate their long-term biocompatibility. In this study, male ICR mice (6–8 weeks old, n = 2) were anesthetized with an intraperitoneal injection of ketamine (100 mg / kg) and xylazine (10 mg / kg) and fixed in a stereotaxic apparatus. As shown in Figure 30, a craniotomy was performed between bregma and lambda (AP -1.5 to -4.5; ML 0.5 to 3.5) on the skull. Three stainless steel bone screws were placed in the skull around the cranial opening, and the dura was then removed. Electrodes were placed on the left and right cortical surfaces, respectively. The cranial opening was filled with Kwik-Sil silicone elastomer (World Precision Instruments, Inc.), and then coated with dental cement (Megadental, Germany) after the Kwik-Sil cured. Mice were sacrificed at 2 weeks and immunohistological analysis was performed to assess their biocompatibility.

[0119] Mice were anesthetized by intraperitoneal injection of ketamine (100 mg / kg) and xylazine (10 mg / kg) and then transcardially perfused with PBS followed by 4% paraformaldehyde (PFA). Brain samples were incubated overnight in PFA and then in 30% sucrose solution for 3 days. Brains were cryosectioned at 30 μm thickness. Slices were stained with Iba1 antibody (Abcam, ab178846) and DAPI (nuclear marker) and observed under a laser confocal microscope (Nikon A1HD25, Japan). Regions of interest for assessing the intensity and soma size of Iba1-positive cells were selected from the tissue located 600 μm below the electrode-brain interface (3 slices per mouse). The intensity and soma size of Iba1-positive cells were analyzed using ImageJ software.

[0120] To investigate the long-term biocompatibility of various electrodes within the biceps femoris muscle, healthy male Sprague-Dawley rats (4–5 weeks old, 200 g) were used for implantation into the biceps femoris muscle. The rats were first treated with light gas anesthesia (isoflurane) and then injected with ketamine (1 wt%) for deep anesthesia. Next, a wound (6 mm wide, 2 mm deep) was created, and electrode arrays (µLME, Au / PI, and Au / PDMS) were attached to the surface of the biceps femoris muscle. After suturing and 1 week of implantation, samples were collected and cryosectioned (25 µm thick) using a CryoStar NX70 Cryostat (ThermoFisher). Slices were stained with H&E (Abcam) and then observed under an inverted microscope (Nikon Ti2-A).

[0121] Hydrogel flexible substrate Fabrication of LM micro-iontronics. Fabrication of the ink on the donor substrate involved several steps. First, a Si wafer was sonicated in acetone, 2-propanol, and deionized (DI) water for 5 minutes each, and then dried with a nitrogen gas blow. Subsequently, a 10 wt% aqueous dextran solution was spin-coated onto the Si wafer at a speed of 4000 rpm for 40 seconds. The wafer was then baked on a hotplate at 80°C for 1 minute and then at 180°C for 30 minutes. After the baking process, the dextran film was spin-coated with negative photoresist NR9-1500P at 4000 rpm for 40 seconds, and then pre-baked on a hotplate at 155°C for 1 minute. The photoresist was applied at 170 mJ / cm using a mask aligner (Suss MA6). 2The photoresist was then post-baked at 105 °C for 3 minutes and developed in DR6 developer for 10-15 seconds. After rinsing the sample in DI water and drying it with compressed nitrogen gas, a 5 nm thick adhesion Cr layer and a 400 nm thick Ag layer were deposited via thermal evaporation, followed by stripping in acetone. Preparation of the hydrogel stamp / substrate involved dissolving a precursor solution consisting of acrylic acid (AA, 30.0 wt%), 2-hydroxy-2-methylpropiophenone (HMPP, 0.4 wt%), and N,N-methylenebisacrylamide (MBAA, 0.5 wt%) in a mixture of water and glycerin (50 wt%). The solution was then exposed to UV light at a wavelength of 365 nm and a power of 15 W for approximately 60 seconds, resulting in the transfer of the Ag electrodes from the Si wafer to the hydrogel.

[0122] Characterization. The crosslink density, relative to the UV exposure duration, was evaluated using the shear storage modulus as a quantitative indicator. Measurements were performed using a TwinDrive rheometer (MCR702; Anton Parr) equipped with a 25 mm diameter cone-plate (CP25) and operated at 1% amplitude and 1 Hz angular frequency. UV exposure began at 30 s and ended at 150 s, and data acquisition was performed using RheoCompass software (Anton Parr). To evaluate the adhesion energy, adhesive samples with dimensions of 50 mm × 25 mm × 2 mm were peeled from both the tissue surface and the device encapsulation / substrate material (50 mm × 25 mm; different materials have different thicknesses). The morphology and surface roughness of the sacrificial layer were characterized using an atomic force microscope (AFM, XE70). The linewidth and thickness of the μLMH were determined using a 3D optical surface profiler (New Zygo NexView). Additionally, a scanning electron microscope (SEM, TESCAN VEGA3) was employed to capture the surface morphology and elemental mapping of the samples. Air permeability tests were performed according to the ASTM D737-08 standard using a MO21S air permeability tester (SDL Americ, Inc.). The airflow pressure was set at 100 kPa. Similarly, moisture permeability tests were performed based on the cup method specified in the E96 / E96M-13 standard. The duration of both the air permeability test and the moisture permeability test was 48 hours, and the temperature and humidity were maintained at constant levels of 22°C and 63%, respectively. The electrical conductivity (σ) of the μLMH was determined by calculating σ = L / (A × R), where L, A, and R represent the length, cross-sectional area, and electrical resistance of the μLMH, respectively. The cross-sectional area (A) of the μLMH was measured using a surface profiler (New Zygo NexView), and the integrated polygon area was calculated using Origin software. The electrical resistance (R) of μLMHs (L = 10 mm) with different line widths was measured using direct current (DC) characterization with a parameter analyzer (Keithley 4200A-SCS parameter analyzer) coupled to a probe station (Micromanipulator).In addition, the electrical resistance of the μLME under different strain levels was evaluated using a four-probe method with a source meter (Keithley 2400) coupled to a customized stretcher (Precise). Furthermore, electrochemical impedance spectroscopy (EIS) was performed using an electrochemical workstation (CHI 660E) equipped with a three-electrode testing system. The reference, counter, and working electrodes were Ag / AgCl (2 M KCl), platinum foil, and μLMH, respectively. The electrolyte consisted of 1x phosphate-buffered saline (PBS) solution (Gibco). Impedance and phase angle data were recorded as a function of frequency in the range of 0.01 Hz to 10,000 Hz using an AC sine wave signal amplitude of 10 mV. EIS spectra of μLMH subjected to various strain levels were acquired using the same electrochemical workstation coupled to a custom stretcher.

[0123] Transcutaneous electrical stimulation and recording of corresponding electromyogram (EMG) signals. All procedures involving epicutaneous application of μLMH patches on the human body complied with ethical guidelines approved by The Hong Kong Polytechnic University (HSEARS20230101001). The electrical stimulation process was generated using an AD Instruments PowerLab 16 / 35. Output voltage and current data were then measured at a sampling frequency of 10 kHz using a DAQ (data acquisition) multimeter system (Keithley DAQ6510). EMG signals were measured at a sampling rate of 1 kHz using a high-precision data acquisition system (AD Instruments PowerLab 16 / 35) and a biological signal amplifier (AD Instruments BioAmp FE132). Raw signal data were digitally filtered using a 50 Hz notch filter followed by a low-pass filter with a cutoff frequency of 30 Hz.

[0124] Rat Surgery. All procedures performed according to the experimental protocol during the animal study were reviewed and approved (approval number A-0664) by the Animal Research Ethics Sub-Committee of the Research Committee at the City University of Hong Kong. Male Sprague-Dawley rats, 7-8 weeks old and weighing 200-300 g, were obtained from the Laboratory Animal Services Center at the Chinese University of Hong Kong. Rats were housed in the Laboratory Animal Research Unit at the City University of Hong Kong under a 12-h light / dark cycle until the experiment. Prior to surgery, rats were first anesthetized with gaseous isoflurane (air mixed with 3% isoflurane at 2 L / min), followed by deep anesthesia with sodium pentobarbital (30 mg / kg). After administration of the anesthetic, the rat head was shaved under controlled temperature conditions at 37°C. An incision was made along the midline of the skin at the craniovertebral junction. The overlying muscles were then cut and retracted to expose the skull. A craniotomy was performed using a dental drill, and the dura mater was carefully removed using fine forceps to ensure complete exposure of the cerebral cortex.

[0125] Viral injection, optogenetic stimulation, and electroencephalogram (ECoG) recording. Sprague-Dawley rats were anesthetized with pentobarbital (50 mg / kg, i.p.) and fixed in a stereotaxic apparatus. Following craniotomy, CaMkii-ChR2-eYFP vector (titer 4 × 10) was injected into the rat. 12A 0.5 μl volume of AAV5 containing 1000kJ of 1000kcal (Taitool) was injected into the somatosensory cortex (coordinates: AP -3.3 mm, ML 3.0 mm, DV 2.0 mm) at a rate of 0.2 μL / min using a 10 μL microsyringe and a 33-gauge metal needle (Hamilton, NV, USA). The injection procedure was controlled by a microsyringe pump (World Precision Instruments, FL, USA) and its controller (WPI). After injection, the needle was held in place for an additional 5 minutes to promote viral diffusion and then slowly withdrawn. The rats were maintained for 3 weeks after viral injection to allow for gene expression. After 3 weeks, the skull was exposed under anesthesia, and a small hole (8 mm wide, 10 mm long) was drilled above the cortex for electrode placement. Two stainless steel bone screws were inserted into the skull around the surgical opening, and the dura mater was removed. A μLMH electrode array was then placed on the cortex. Finally, a low-resistance 200 μm silver wire connected to each array was wrapped around one of the bone fixation screws for grounding purposes. In the optogenetic manipulation group, each optical fiber was connected to a blue light laser (MBL-473 / 200 mW; Fiblaser) with an output of 7–10 mW / mm. 2 473 nm blue light was delivered to each fiber optic terminal within the somatosensory cortical region at a frequency of 20 Hz (square wave with a 10% duty cycle). The μLMH array, coupled to a flexible printed circuit board (FPC) connector, was connected to the PCB using a customized FPC adapter. The resulting ECoG neural signals were recorded using a multichannel high-precision data acquisition system (AD Instruments PowerLab 16-35). The signals were acquired and digitally bandpass filtered (1–30 Hz) using LabChart software. Custom-written Matlab code was used to analyze the ECoG neural signal spectrograms.

[0126] Biocompatibility testing of μLMH neural implants. After a 2-week neural implantation period, mice were sacrificed for subsequent immunohistological analysis. To ensure anesthesia, ketamine (100 mg / kg) and xylazine (10 mg / kg) were administered intraperitoneally. Mice were then transcardially perfused with PBS followed by 4% paraformaldehyde (PFA). Brain samples were then incubated overnight in PFA and in 30% sucrose solution for 3 days. Brain cryosections were performed at 30 μm thickness. Slices were stained with Iba1 antibody (Abcam, ab178846) and DAPI (a nuclear marker) and examined using a laser confocal microscope (Nikon A1HD25, Japan). Regions of interest, including the electrode-brain interface and tissue 600 μm below the interface, were selected to evaluate the intensity and soma size of Iba1-positive cells (three slices were analyzed per mouse). ImageJ software was employed to analyze both the intensity and soma size of Iba1-positive cells. [Explanation of symbols]

[0127] 100 Bioelectronic Devices 101 Flexible substrate 102 Electronic Components 103 Patterned Metal Structures 104 Sacrificial Layer 105 PCB

Claims

1. 1. A permeable stretchable bioelectronic device comprising: a flexible substrate; and one or more electronic components disposed on a surface of the flexible substrate, wherein the flexible substrate comprises a biocompatible elastomeric fiber or a biocompatible hydrogel, and wherein each of the one or more electronic components independently comprises a metal and a liquid metal.

2. 10. The permeable stretchable bioelectronic device of claim 1, wherein each of the one or more electronic components is independently selected from the group consisting of a transistor, a diode, an electrode, a generator, an inductor, a capacitor, and a resistor.

3. 10. The permeable stretchable bioelectronic device of claim 1, wherein the biocompatible elastomeric fiber is selected from the group consisting of an elastomeric homopolymer, an elastomeric block copolymer, an elastomeric random copolymer, an elastomeric graft copolymer, an elastomeric brush copolymer, an elastomeric thermoset, and combinations thereof; and the biocompatible hydrogel comprises a hydrophilic polymer selected from the group consisting of poly(acrylic acid), poly(vinyl alcohol), poly(ethylene oxide), poly(ethylene glycol), gelatin, collagen, polyacrylamide, polysaccharides, and combinations thereof, wherein the hydrophilic polymer is optionally crosslinked.

4. 10. The permeable stretchable bioelectronic device of claim 1, wherein the biocompatible elastomeric fiber comprises styrene-isoprene-styrene block copolymer, styrene-polybutadiene-styrene block copolymer, styrene-butadiene block copolymer, poly(styrene-block-butadiene-block-styrene) copolymer, polyisoprene rubber, butadiene rubber, or mixtures thereof, and the biocompatible hydrogel comprises crosslinked poly(acrylic acid).

5. 10. The permeable stretchable bioelectronic device of claim 1, wherein the biocompatible elastomeric fiber comprises a styrene-polybutadiene-styrene block copolymer and the biocompatible hydrogel comprises poly(acrylic acid) crosslinked with N,N'-methylenebisacrylamide.

6. 10. The permeable stretchable bioelectronic device of claim 1, wherein the biocompatible elastomeric fibers have an average diameter of 0.1 μm to 100 μm.

7. 10. The permeable stretchable bioelectronic device of claim 1, wherein the flexible substrate comprising the biocompatible elastomeric fibers is plasma treated.

8. 10. The permeable stretchable bioelectronic device of claim 1, wherein the flexible substrate has a thickness of 25 μm to 330 μm.

9. 10. The permeable stretchable bioelectronic device of claim 1, wherein the metal comprises copper, silver, gold, or a mixture thereof.

10. 10. The permeable stretchable bioelectronic device of claim 1, wherein the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

11. 10. The permeable stretchable bioelectronic device of claim 1, wherein the liquid metal comprises a eutectic gallium indium alloy (EGaln), a gallium indium tin alloy (GaInSn), or a mixture thereof.

12. The one or more electronic components are 100 to 75,000 electronic components / cm ２ 10. The permeable stretchable bioelectronic device of claim 1, wherein the nanoparticles are present on the surface of the flexible substrate at a density of

13. 10. The permeable stretchable bioelectronic device of claim 1, wherein the flexible substrate comprises poly(styrene-block-butadiene-block-styrene) copolymer or poly(acrylic acid) crosslinked with N,N'-methylenebisacrylamide, the one or more electronic components comprise silver and eutectic gallium indium alloy (EGaln), and the flexible substrate comprising poly(styrene-block-butadiene-block-styrene) copolymer is plasma treated.

14. 14. The permeable stretchable bioelectronic device of claim 13, wherein the one or more electronic components include an electrode.

15. 10. A method of making the permeable stretchable bioelectronic device of claim 1, comprising: providing a substrate having a sacrificial layer formed thereon, said sacrificial layer comprising a water-soluble material selected from the group consisting of water-soluble polymers, water-soluble sugars, water-soluble metal salts, and mixtures thereof; depositing a photoresist layer on the sacrificial layer; irradiating the photoresist layer using a patterned mask, thereby forming an irradiated photoresist layer including an exposed photoresist layer and an unexposed photoresist layer; contacting the irradiated photoresist layer with a developer, thereby removing either the exposed or unexposed photoresist layer and forming a patterned photoresist layer including an exposed patterned sacrificial layer; depositing a metal on the exposed and patterned sacrificial layer; removing the irradiated photoresist layer, thereby forming one or more patterned metal structures on the sacrificial layer; depositing the biocompatible elastomeric fiber onto the one or more patterned metal structures, thereby forming a flexible substrate comprising one or more patterned metal structures disposed on a surface of the flexible substrate; or depositing a hydrogel precursor solution onto the one or more patterned metal structures, thereby forming a hydrogel precursor coating, and curing the hydrogel precursor coating, thereby forming the flexible substrate comprising one or more patterned metal structures disposed on a surface of the flexible substrate; contacting the sacrificial layer with an aqueous solvent, thereby at least partially dissolving the sacrificial layer; separating the substrate and the flexible substrate, the flexible substrate including one or more patterned metal structures disposed on a surface of the flexible substrate; depositing the liquid metal onto the one or more patterned metal structures, thereby forming the permeable stretchable bioelectronic device; A method comprising:

16. The method of claim 15 , wherein the biocompatible elastomeric fibers are deposited using an electrospinning process.

17. The method of claim 15, wherein the flexible substrate including the biocompatible elastomeric fiber is subjected to a plasma treatment for 1 to 30 minutes.

18. 16. The method of claim 15, wherein the photoresist layer comprises a negative photoresist composition, and the developer removes the unexposed photoresist layer.

19. 16. The method of claim 15, wherein the metal comprises copper, silver, gold, or a mixture thereof, and the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

20. 16. The method of claim 15, wherein the biocompatible elastomeric fiber is deposited using an electrospinning method, the flexible substrate is subjected to a plasma treatment for 1 to 30 minutes, the metal comprises copper, silver, gold, or a mixture thereof, and the liquid metal comprises gallium, a gallium alloy, or a mixture thereof.

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