Waterproof, GEL-free, wireless and wearable multiple channel bioelectric recording system and methods of use thereof

Abstract

A waterproof, gel-free, wireless and wearable multiple channel bioelectric recording system that includes E-textile bioelectric sensors, methods of fabricating the E-textile bioelectric sensors, and methods of monitoring electrophysiological signals such as ECG and EMG signals using the bioelectric recording system are disclosed.

Description

TITLE OF THE INVENTIONWATERPROOF, GEL-FREE, WIRELESS AND WEARABLE MULTIPLE CHANNEL BIOELECTRIC RECORDING SYSTEM AND METHODS OF USE THEREOFCROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims the benefit of priority to US Provisional Application No. 63 / 595,655 filed on 02 November 2023, the content of which is incorporated by reference herein in its entirety.STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

[0002] This invention was made with government support under HD104822 and HD105905 awarded by the National Institutes of Health. The government has certain rights in the invention.MATERIAL INCORPORATED-BY-REFERENCE

[0003] Not applicable.FIELD OF THE INVENTION

[0004] The present disclosure generally relates to a wearable electrodes, bioelectric recording systems, and methods of use thereof.BACKGROUND OF THE INVENTION

[0005] The rapid advancement of soft wearable electronics is propelling a revolution in the field of activity tracking, health monitoring and clinical medical devices by providing greatly enhanced user experience and patient compliance. Although sensor patches that are attached on skin or sensors and electronics that are directly fabricated on skin are widely recognized as effective approaches, they may lack the desired unobtrusiveness and require complicated fabrication processes, thus preventing them from being ready-to- use and comfortable for long-term wearability. Textiles have emerged as a platform for seamlessly integrating various smart functionalities, including sensing, energy harvesting, displaying, and wireless communication. Diversematerials have been explored for constructing conductors on textiles, including metals, carbon-based materials, and conjugated polymers. Mainstream electronic textile (E-textile) manufacturing methods involve knitting / weaving / embroidering one-dimensional (1 D) fiber / yarn-shaped conductive patterns into finished textile, or soaking / coating / printing conductive ink / paste onto fabric surface to create 2D conductive elements. Especially, screen printing stands out as a promising approach due to its fast, cost- effective, and scalable patterning capability for mass production.

[0006] Real-time monitoring of electrophysiological signals is crucial in medical diagnosis, health monitoring, and exercise physiology studies. While E-textiles have great potential for applications in those scenarios, there are also a few key challenges, including maintaining a reliable skin-electrode contact for high- quality signal recording, mitigating motions artifacts from ambulatory subjects, and the necessity to eliminate the need for conductive hydrogel for long-term usability. When compared to conventional gel-assisted planar electrodes, gel- free dry electrodes hold great potential in addressing issues such as skin irritation, discomfort caused by tightening / pressing operation for improved contact, and signal degradation over time caused by gel dehydration.

[0007] Additionally, the practical application of textile-based electronics is hindered by issues of durability and reliability. For instance, perspiration and ambient moisture can lead to electrode erosion or short circuits, resulting in performance degradation. Polymer encapsulation or lamination have been validated as effective approaches for packaging the devices, but the use of polymer layers coated on textile compromises the natural permeability, breathability, and softness of the fabrics. Lastly, the feasibility of adopting E- textiles in real-life applications can also be influenced by manufacturing cost and consumer expenses for usage and maintenance.SUMMARY OF THE INVENTION

[0008] Among the various aspects of the present disclosure is the provision of a waterproof, gel-free, wireless and wearable multiple channel bioelectric recording system.

[0009] Briefly, therefore, the present disclosure is directed to a bioelectric recording system and methods of use thereof.

[0010] The present teachings include a wearable, wireless E-textile bioelectric recording system. In one aspect, the system can include at least two 3D soft and fluffy microfiber-based electrodes printed on a textile to greatly increase the surface area. In another aspect, the system can include a superhydrophobic fluorinated self-assembled monolayer deposited on the electrode-containing E-textile surface to render it waterproof while retaining the electrical conductivity. In yet another aspect, the system can include a custom- designed motion-artifact cancelling wireless data recording circuit to detect bioelectric signals from the electrodes. In some aspects, the soft and fluffy conductive microfibers can disperse freely and securely adhere to the skin of a subject, achieving a low impedance at the electrode-skin interface even in the absence of gel. In accordance with another aspect, the microfiber-based electrodes can be composed of a conducting polymer. In one aspect, the conducting polymer can be PEDOT:PSS. In another aspect, the selfassembled monolayer can be composed of perfluorooctyltrichlorosilane (PFOTS). In yet another aspect, the system can be equipped with at least two recording channels. In some embodiments, the at least 2 recording channels record at least one electrocardiogram (ECG) and at least one electromyogram (EMG) simultaneously. In some embodiments, the system can incorporate spatial-temporal mapping. In yet another aspect, the system can be integrated into a smart garment to be worn by a subject.

[0011] The present teachings also include a method to monitor the health and physiology of a subject. In one aspect, the method can include having the subject wear a smart garment containing the E-textile bioelectric recording system as described herein on or in a region of interest and recording bioelectric signals of the subject with the E-textile system. In another aspect, the system can perform real-time multimodal electrophysiological signal monitoring. In another aspect, the real-time signal monitoring can be selected from an electrocardiogram (ECG), an electromyography (EMG), and any combination thereof. In some embodiments, the real-time monitoring can becarried out during strenuous exercise by a subject. In other embodiments, the ECG can be a maternal ECG, and the EMG can be a uterine EMG of a pregnant mother. In another aspect, the maternal ECG and the uterine EMG contributes to uterus-relevant physiological assessments and early detection of abnormal uterine contraction patterns and birth-related risks.

[0012] The present teaching also includes a method to fabricate an E-textile bioelectric recording device. In one aspect, the method can include printing at least two 3D microfiber-based electrodes onto a textile. In another aspect, the method can include vaporizing a self-assembled monolayer of a superhydrophobic monolayer onto the electrodes. In yet another aspect, the method can include connecting a custom-designed motion-artifact cancelling wireless data recording circuit to the electrodes to record bioelectric signals. In accordance with another aspect, the electrodes can be composed of a conducting polymer. In another aspect, the conducting polymer can be PEDOT:PSS. In yet another aspect, the superhydrophobic monolayer is composed of perfluorooctyltrichlorosilane (PFOTS). In other aspects, the device is equipped with at least two recording channels by connecting at least two electrodes to the recording circuit.

[0013] Other objects and features will be in part apparent and in part pointed out hereinafter.DESCRIPTION OF THE DRAWINGS

[0014] The following drawings illustrate various aspects of the disclosure.

[0015] FIG. 1 is a schematic drawing of a waterproof and gel-free 3D microfiber-on-textile electrode in accordance with an aspect of the disclosure.

[0016] FIG. 2A is a top-view SEM image of the textile interface between the PEDOT:PSS-coated (left) and non-coated (right) regions. Scale bar, 500 pm.

[0017] FIG. 2B is a schematic showing the fibers in a textile and the current flow path.

[0018] FIG. 2C is a graph summarizing the evolution of sheet resistance for on- textile electrodes during multiple stretching tests of 5%, 15%, and 25% insequence.

[0019] FIG. 2D is a schematic diagram illustrating the vapor deposition of PFOTS to achieve a waterproof E-textile.

[0020] FIG. 2E is a graph illustrating the evolution of droplet contact angles on PFOTS-treated textile over 7 days, in comparison with the contact angle on the untreated textile and the PFOTS-treated one submerged in water for an hour. The inset images show the microscopic images of water droplets on the corresponding textile. Scale bar, 500 pm.

[0021] FIG. 2F is a graph summarizing the sheet resistance of PFOTS-treated on-textile electrodes during the process of submerging in water, bending, 180° twisting, 360° twisting, and redrying. The insets show the photographs of water droplets balling up on the as-prepared PFOTS-treated textile, along with the corresponding operations under water. Scale bar, 1 cm.

[0022] FIG. 2G is an image of a waterproof and gel-free electrode that includes an on-textile base electrode, an ionic binder, and an add-on conductive microfiber layer.

[0023] FIG. 2H is an SEM image of the non-treated microfibers. Scale bars, 50 pm.

[0024] FIG. 2I is an SEM image of the PEDOT:PSS-coated conductive microfibers. Scale bars, 25 pm.

[0025] FIG. 2J is a graph summarizing the electrode-skin contact impedance of several electrode designs over the frequency range of 10 to 1000 Hz under dry conditions. The inset image is a photograph of a functional single electrode made by adding conductive microfibers onto the on-textile base electrode. Scale bar, 1 cm.

[0026] FIG. 3A is a graph showing ECG recordings obtained using the on- textile electrodes under dry (upper graph) and sweating (lower trace) conditions.

[0027] FIG. 3B is a graph showing ECG recordings obtained using the microfiber-on-textile electrodes under dry (upper graph) and sweating (lowertrace) conditions.

[0028] FIG. 3C is a graph comparing SNR measured using between using on- textile electrodes and microfiber-on-textile electrodes.

[0029] FIG. 3D contains a series of ECG traces obtained from on-textile electrodes after multiple washing-redry cycles.

[0030] FIG. 3E is a graph summarizing the SNRs obtained from on-textile electrodes after multiple washing-redry cycles. The inset image illustrates the experimental implementation of the washing test, including agitation in warm water (40 °C) and drying process.

[0031] FIG. 3F contains a series of ECG traces obtained from microfiber-on- textile electrodes after multiple washing-redry cycles over a period of up to 4 months.

[0032] FIG. 3G is a graph summarizing the SNRs obtained from on-textile electrodes after multiple washing-redry cycles over a period of up to 4 months.

[0033] FIG. 3H contains EMG recordings obtained using the E-textile equipped with microfiber-on-textile electrodes under dry conditions by lifting various weights.

[0034] FIG. 3I contains EMG recordings obtained using the E-textile equipped with microfiber-on-textile electrodes under sweating conditions by lifting various weights.

[0035] FIG. 3J is a graph of the corresponding SNR values of the EMG recordings od FIGS. 3H and 3I.

[0036] FIG. 4A is a block diagram of the portable instrumentation system for recording electrophysiological signals.

[0037] FIG. 4B contains graphs of raw ECG signals (center trace) and heart rate analysis (lower trace) obtained over a 14-m inute cycling session including the warm-up, climbing, and sprinting phases. Zoomed-in views of the ECG signals within 6-second windows during different training phases are provided in the upper row of graphs.

[0038] FIG. 4C contains a series of graphs summarizing real-time EMG monitoring and muscle output analysis of a subject’s quadricep during a 6- minute cycling session, including the warm-up, climbing, sprinting, and cooldown phases, including recorded raw thigh-EMG signals (upper trace) with zoomed-in views of the EMG signals within 10-second windows during different training phases, training information from the stationary bike including cadence (upper center trace) and resistance (upper bottom trace), and real-time muscle output analysis from the E-textile system including cadence (lower center trace) and EMG amplitude (lower bottom trace).

[0039] FIG. 5A contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while walking to the pool from the locker room.

[0040] FIG. 5B contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while stepping into the pool.

[0041] FIG. 5C contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while floating and diving.

[0042] FIG. 5D contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while swimming.

[0043] FIG. 5E contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while stepping out of the pool.

[0044] FIG. 5F contains a real-time raw ECG recording obtained using the E- textile system from a swimmer during a workout while walking back to the locker room.

[0045] FIG. 6A a system-level block diagram of the multi-channel E-textile system, including the multi-channel signal transduction, signal processing and conditioning, wireless data transmission to the portable electronic device, channel demultiplexing, and corresponding post-processing to display the readable multi-channel ECG and multi-channel EMG signals.

[0046] FIG. 6B is a schematic illustration depicting the E-textile patch equipped with the multi-channel data recording electronics, designed for real- timematernal health monitoring in clinical settings by attaching the system to the subject’s abdomen.

[0047] FIG. 6C contains graphs related to real-time maternal ECG monitoring and data analysis (e.g. real-time heart rate and time-frequency analysis) of a pregnant subject including raw maternal ECG signals with extracted excerpts (top trace), the extracted heart rate (center trace), and the time-frequency analysis (bottom map).

[0048] FIG. 6D contains graphs related to real-time uterine EMG monitoring and RMS analysis from the wireless E-textile system (top trace pair), in comparison with the recordings from commercial BioSemi system (center trace) and the tocodynamometer (bottom trace).

[0049] FIG. 7A contains readings from a representative channel (#4) of a multichannel E-textile system obtained during real-time uterine EMG monitoring including the EMG-RMS waveforms spanning 67 minutes.

[0050] FIG. 7B is a zoomed-in view of contraction #20 from FIG. 7A as captured by multiple channels of the E-textile system

[0051] FIG. 7C contains graphs illustrating the evolution of uterine contraction intervals (lower graph) and the analysis of monitoring accuracy (upper graph) using the E-textile system in comparison to the TOCO system. The linear fitting yy = kkkk + bb was performed to exhibit the trend of interval evolutions along the labor progression (ii), with similar slope value fcfcl = -6.09 (line, E- textile) and kk2 = -6.14 (dashed line, TOCO).

[0052] FIG. 8A is a block diagram illustrating a process of data processing for the multiplexed inputs from a multi-channel E-textile system used to obtain both multi-channel maternal ECG and uterine EMG outputs.

[0053] FIG. 8B is a graph showing a power spectral density obtained from fast Fourier transform (FFT) according to the process illustrated in FIG. 8A.

[0054] FIG. 8C is a graph of data from a channel from a multi-channel E-textile system filtered to isolate the maternal ECG portion of the signals.

[0055] FIG. 8D is a graph of data from a channel from a multi-channel E-textilesystem filtered to isolate the uterine EMG portion of the signals.

[0056] FIG. 9A a 20-second extract of ECG signals collected by an E-textile system as part of a clinical study of real-time maternal health monitoring.

[0057] FIG. 9B is series of graphs summarizing a 15-second extract of EMG signals collected by an E-textile system (top graph), including an RMS analysis (upper center graph) as well as comparison signals obtained using a commercial BioSemi system (lower center graph) and a tocodynamometer (TOCO, bottom graph).

[0058] FIG. 10A is a 20-second ECG trace collected by the E-textile system.

[0059] FIG. 10B contains a 15-second extract of uterine EMG signals collected by an E-textile system (top graph), including an RMS analysis (upper center graph) as well as comparison signals obtained using a commercial BioSemi system (lower center graph) and an intrauterine pressure catheter (IUPC).

[0060] FIG. 11 A is a 20-second ECG trace collected by the E-textile system.

[0061] FIG. 11 B contains a 15-second extract of uterine EMG signals collected by an E-textile system (top graph), including an RMS analysis (upper center graph) as well as comparison signals obtained using a commercial BioSemi system (lower center graph) and a tocodynamometer (TOCO, bottom graph) that failed due to the wrong placement induced by the subject’s motion.

[0062] FIG. 12A is a chemical structure diagram of PEDOT:PSS and PEO, and step-by-step reactions of PEO with PSS.

[0063] FIG. 12B is a graph summarizing sheet resistance of printed electrodes on textiles under varying strain levels using PEDOT:PSS inks, with and without PEO additive.

[0064] FIG. 12C is a graph summarizing the evolution of sheet resistance for on-textile PEDOT:PSS (without PEO) electrodes during multiple stretching tests of 5%, 15%, and 25% in sequence.

[0065] FIG. 13A contains an SEM image of a pristine textile.

[0066] FIG. 13B is an EDS scanning result obtained within the circled region ofFIG. 13A.

[0067] FIG. 13C is a graph of EDS elemental analyzing results obtained within the circled region of FIG. 13A.

[0068] FIG. 13D contains an SEM image of a PEDOT:PSS-coated textile.

[0069] FIG. 13E is an EDS scanning result obtained within a region of FIG. 13D.

[0070] FIG. 13F is a graph of EDS elemental analyzing results obtained within the region of FIG. 13D.

[0071] FIG. 14A contains a current-flowing model and stretchability characterizations (5%, 15%, and 25% in sequence) of the PEDOT:PSS conductors printed on different sides of nylon-spandex knitted fabric along a wale direction on the front side. The stretching operations were repeated seven times at each state. Scale bars, 500 pm.

[0072] FIG. 14B contains a current-flowing model and stretchability characterizations (5%, 15%, and 25% in sequence) of the PEDOT:PSS conductors printed on different sides of nylon-spandex knitted fabric along a course direction on the front side. The stretching operations were repeated seven times at each state. Scale bars, 500 pm.

[0073] FIG. 14C contains a current-flowing model and stretchability characterizations (5%, 15%, and 25% in sequence) of the PEDOT:PSS conductors printed on different sides of nylon-spandex knitted fabric along a wale direction on the back side. The stretching operations were repeated seven times at each state. Scale bars, 500 pm.

[0074] FIG. 14D contains a current-flowing model and stretchability characterizations (5%, 15%, and 25% in sequence) of the PEDOT:PSS conductors printed on different sides of nylon-spandex knitted fabric along a course direction on the back side. The stretching operations were repeated seven times at each state. Scale bars, 500 pm.

[0075] FIG. 15A is a schematic diagram showing the functionalization of the textile by vaporizing 1 H,1 H,2H,2H-perfluorooctyltrichlorosilane.

[0076] FIG. 15B is a schematic diagram showing and the self-assembled monolayer (SAM) on the surface of the textile.

[0077] FIG. 15C is a top-view SEM image of the pristine (nylon-spandex textile used to build the E-textile system. Scale bar, 500 pm.

[0078] FIG. 15D is a top-view SEM image of a heavily functionalized nylon- spandex textile used to build the E-textile system. Scale bar, 500 pm.

[0079] FIG. 16A is a photograph of water drops on a cotton textile waterproofed as illustrated in FIG 15A.

[0080] FIG. 16B is a photograph of water drops on a polyester textile waterproofed as illustrated in FIG 15A.

[0081] FIG. 16C is a photograph of water drops on a cleanroom cloth textile waterproofed as illustrated in FIG 15A.

[0082] FIG. 16D is a photograph of water drops on a cotton textile waterproofed as illustrated in FIG 15A.

[0083] FIG. 17 is a graph summarizing the impedance between an on-textile electrode and skin under dry, sweating, and gel conditions.

[0084] FIG. 18A contains a series of ECG waveforms recorded using conductive microfiber electrodes after various numbers of washing cycles and drying in oven.

[0085] FIG. 18B contains a series of ECG waveforms recorded using conductive microfiber electrodes after various numbers of washing cycles and drying by a hair dryer.

[0086] FIG. 18C is a graph comparing the SNRs of the ECG waveforms recorded using conductive microfiber electrodes subjected to the washing & drying tests of FIGS. 18A and 18B.

[0087] FIG. 19A contains an SEM image of PEDOT:PSS-coated conductive microfibers dried in an oven.

[0088] FIG. 19B contains an enlarged SEM image of the PEDOT:PSS-coated conductive microfibers of FIG. 19A.

[0089] FIG. 19C contains an SEM image of PEDOT:PSS-coated conductive microfibers dried by hair dryer.

[0090] FIG. 19D contains an enlarged SEM image of the PEDOT:PSS-coated conductive microfibers of FIG. 19D.

[0091] FIG. 19E contains an SEM image of the PEDOT:PSS-coated conductive microfibers of FIG. 19A rewashed and redried in an oven.

[0092] FIG. 19F contains an enlarged SEM image of the PEDOT:PSS-coated conductive microfibers of FIG. 19E.

[0093] FIG. 19G contains an SEM image of PEDOT:PSS-coated conductive microfibers of FIG. 19C rewashed and redried by hair dryer.

[0094] FIG. 19H contains an enlarged SEM image of the PEDOT:PSS-coated conductive microfibers of FIG. 19G.

[0095] FIG. 20A contains graphs summarizing impedance between skin and conductive microfibers coated with PEO-free PEDOT:PSS ink.

[0096] FIG. 20B is contains graphs summarizing impedances between skin and conductive microfibers coated with PEDOT:PSS / PEO composite ink in a 10 / 1 weight ratio.

[0097] FIG. 20C contains graphs summarizing impedance between skin and conductive microfibers coated with PEDOT:PSS / PEO composite ink in a 5 / 1 weight ratio.

[0098] FIG. 20D contains graphs summarizing ECG signals collected by the conductive microfibers coated with PEO-free PEDOT:PSS ink.

[0099] FIG. 20E contains graphs summarizing ECG signals collected by the conductive microfibers coated with PEDOT:PSS / PEO composite ink in a 10 / 1 weight ratio.

[0100] FIG. 20F contains graphs ECG signals collected by the conductive microfibers coated with PEDOT:PSS / PEO composite ink in a 5 / 1 weight ratio.

[0101] FIG. 21 A contains graphs summarizing the evolution of ECG signals recorded by conductive microfibers coated with 10 / 1 PEDOT:PSS / PEOcomposite ink over multiple washing and drying cycles.

[0102] FIG. 21 B contains graphs summarizing the evolution of ECG signals recorded by conductive microfibers coated with 5 / 1 PEDOT:PSS / PEO composite ink over multiple washing and drying cycles.

[0103] FIG. 21 C is a graph summarizing changes in SNR signals of FIGS. 21 A and 21 B over the multiple washing / drying cycles.

[0104] FIG. 22 contains an array of SEM images summarizing the degradation of the conductive coating on microfibers after multiple washing / drying cycles for conductive microfibers coated with 10 / 1 PEDOT:PSS / PEO composite ink (left) and coated with 5 / 1 PEDOT:PSS / PEO composite ink.

[0105] FIG. 23A is a schematic of a stitching method of bonding the on-textile electrode and the add-on conductive microfibers.

[0106] FIG. 23B is a graph summarizing the impedance between skin and microfiber-on-textile electrodes stitched with conductive thread using the method illustrated in FIG. 23A. The conductive thread was prepared by dipcoating cotton threads in PEDOT:PSS ink, followed by oven drying and annealing.

[0107] FIG. 23C is a graph summarizing the impedance between skin and microfiber-on-textile electrodes stitched with metal wire using the method illustrated in FIG. 23A.

[0108] FIG. 23D is a schematic of a binder-based method of bonding the on- textile electrode and the add-on conductive microfibers.

[0109] FIG. 23E is a graph summarizing the impedance between skin and microfiber-on-textile electrodes bonded with Ag epoxy using the method illustrated in FIG. 23D.

[0110] FIG. 23F is a graph summarizing the impedance between skin and microfiber-on-textile electrodes bonded with carbon tape using the method illustrated in FIG. 23D.

[0111] FIG. 23G is a graph summarizing the impedance between skin andmicrofiber-on-textile electrodes bonded with ionic binder using the method illustrated in FIG. 23D.

[0112] FIG. 24 is a graph summarizing an EDS element analysis result of conductive microfibers coated with 5 / 1 PEDOT:PSS / PEO composite ink.

[0113] FIG. 25A contains a schematic diagram (left) and equivalent circuit model (right) for a conventional gelled electrode on skin.

[0114] FIG. 25B contains a schematic diagram and overlaid equivalent circuit model for a gel-free on-textile electrode on skin.

[0115] FIG. 25C contains a schematic diagram for a gel-free on-textile electrode on skin; the equivalent circuit model is similar to the model illustrated in FIG. 25B.

[0116] FIG. 26A contains graphs summarizing ECG signals recorded by on- textile electrodes under dry, sweating, and gel conditions.

[0117] FIG. 26B contains graphs summarizing ECG signals recorded by microfiber-on-textile electrodes under dry, sweating, and gel conditions.

[0118] FIG. 26C is a graph comparing the SNRs of the ECG signals of FIGS. 26A and 26B.

[0119] FIG. 27A contains graphs of ECG signals collected by pristine on-textile electrodes over multiple washing cycles.

[0120] FIG. 27B is a graph comparing the SNRs of the ECG signals of FIG. 27A.

[0121] FIG. 28A contains a real-time ECG waveform obtained using an on- textile electrode with microfibers added under the presence of motion artifacts caused by periodic body movement.

[0122] FIG. 28B contains a real-time ECG waveform obtained using an on- textile electrode with the base electrode only under the presence of motion artifacts caused by periodic body movement.

[0123] FIG. 28C contains a real-time ECG waveform obtained using commercial Ag / Ag / CI electrodes under the presence of motion artifacts causedby periodic body movement.

[0124] FIG. 29A contains a series of EMG waveforms recorded using commercial Ag / AgCI electrodes.

[0125] FIG. 29B contains a series of EMG waveforms recorded using microfiber-on-textile electrodes.

[0126] FIG. 30 contains a circuit diagram of an analog frontend of a singlechannel sensing system, consisting of filtering network, input common-mode circuits, and the instrumentation amplifier.

[0127] FIG. 31 contains a circuit diagram of a multiplexed analog frontend in the multi-channel sensing system, consisting of differential analog multiplexer, active shielding circuit, filtering network, input common-mode circuits, and the instrumentation amplifier.

[0128] FIG. 32A compares EMG waveforms associated with uterine contraction recorded by E-textile sensors (upper trace) and TOCO systems (lower trace).

[0129] FIG. 32B compares EMG waveforms associated with uterine contraction recorded by E-textile sensors (upper trace) and TOCO systems (lower trace).

[0130] FIG. 33A is a timeline for performing measurements using an E-textile system during clinical patient studies.

[0131] FIG. 33B is a timeline for performing measurements using an existing BioSemi system during clinical patient studies.

[0132] FIG. 34 contains a drawing of a mesh-type mask for screen printing the multi-channel E- textile patch used for clinical patient studies.

[0133] Those of skill in the art will understand that the drawings, described below, are for illustrative purposes only. The drawings are not intended to limit the scope of the present teachings in any way.DETAILED DESCRIPTION OF THE INVENTION

[0134] The present disclosure is based, at least in part, on the recent development of waterproof and gel-free electrodes and wireless analog-to- digital converters that can be used to support a comfortable and wearablesolution of multi-channel bioelectricity recording system. In some aspects, this system can be employed for electromyometrial imaging (EMMI) and Uterine peristalsis Imaging (UPI). As shown herein, waterproof and gel-free 3D E- textile systems for exercise physiology and clinical maternal health monitoring are described.

[0135] One aspect of the present disclosure provides a cost-effective E-textile system (costing as low as $1 per sensor) that features 3D microfiber-based electrodes for greatly increasing the surface area. The soft and fluffy conductive microfibers can disperse freely and securely adhere to the skin, achieving a low impedance at the electrode-skin interface even in the absence of gel. A superhydrophobic fluorinated self-assembled monolayer can be deposited on the E-textile surface to render it waterproof while retaining the electrical conductivity. Equipped with a custom-designed motion-artifact cancelling wireless data recording circuit, the E-textile system can be integrated into a variety of smart garments for exercise physiology and health monitoring applications. Real-time multimodal electrophysiological signal monitoring, including electrocardiogram (ECG) and electromyography (EMG), was successfully carried out during strenuous cycling and even underwater swimming activities. Furthermore, a multi-channel E-textile was developed and implemented in clinical patient studies for simultaneous real-time monitoring of maternal ECG and uterine EMG signals, incorporating spatial-temporal potential mapping capabilities. Such advancement can contribute to uterusrelevant physiological assessments and early detection of abnormal uterine contraction patterns and birth-related risks, thereby improving prenatal care and providing valuable insights for women’s health.

[0136] Electronic textiles (E-textiles) offer great wearing comfort and unobtrusiveness, thus holding potential for next-generation health monitoring wearables. However, the practical implementation is hampered by challenges associated with poor signal quality, substantial motion artifacts, durability for long-term usage, and non-ideal user experience. A cost-effective E-textile system comprising 3D microfiber-based electrodes for greatly increasing the surface area is disclosed. The soft and fluffy conductive microfibers dispersefreely and securely adhere to the skin, achieving a low impedance at the electrode-skin interface even in the absence of gel. In some aspects, a superhydrophobic fluorinated self-assembled monolayer is deposited on the E- textile surface to render it waterproof while retaining the electrical conductivity of the 3D microfiber-based electrodes. In some aspects, the E-textile system further comprises a wireless data recording circuit configured to receive and record sensor measurements form the E-textile sensors as well as process the measured data to remove any motion artifacts caused by movements of the wearer, including but not limited to movements associated with exercise. In various aspects, the E-textile system is suitable for integration into a variety of smart garments for monitoring electrophysiological activity used in exercise physiology and health monitoring applications.

[0137] Win various aspects, an E-textile system for monitoring electrophysiological signals of a subject is disclosed that integrates microfiber- on-textile electrodes and miniaturized circuitry, enabling comfortable and wearable electrophysiology monitoring in a variety of challenging conditions. The free-standing microfibers of the integrates microfiber-on-textile electrodes provide for on-body recordings of multiple bipotential signals in a gel-free and motion-artifact-tolerant manner. The microfiber-on-textile electrodes are functionalized for water-repellence by vapor-depositing self-assembled PFOTS monolayers, imparting the microfiber-on-textile electrodes with sweat-tolerant and waterproof capability and ensuring uninterrupted electrophysiology monitoring even in conditions of severe perspiration or underwater usage.

[0138] In various aspects, the E-textile system includes planar printed on-textile interconnects, add-on conductive microfibers, ionic binders, and mini-PCBs. These features of the E-textile system are readily assembled and integrated into garments of various form factors including, but not limited to athletic wear and wearable clinical monitoring devices. This versatility provides for diverse sports physiology applications in real-life scenarios such as strenuous cycling and water sports. In addition, clinical studies as described in the Examples herein validated the E-textile system's capability for multimodal monitoring of maternal health, including maternal ECG, uterine EMG, and potential maps.

[0139] As described in the examples herein, real-time multimodal electrophysiological signal monitoring, including electrocardiogram (ECG) and electromyography (EMG), was successfully carried out during strenuous cycling and underwater swimming activities. In other examples described herein, a multi-channel E-textile system was developed and implemented in clinical patient studies for simultaneous real-time monitoring of maternal ECG and uterine EMG signals; the multi-channel E-textile system provided for spatial-temporal potential mapping capabilities. The use of the E-textile system for maternal ECG and uterine EMG monitoring provides for uterus-relevant physiological assessments and early detection of abnormal uterine contraction patterns and birth-related risks, thereby improving prenatal care and providing valuable insights for women’s health.

[0140] In various aspects, a 3D microfiber-on-textile electrode 100 is shown illustrated in FIG. 1. The electrode 100 includes an on-textile base electrode 102 formed from a conductive polymer including, but not limited to, a PEDOT:PSS conductive polymer described in additional detail herein. In some aspects, the conductive PEDOT:PSS polymer may further include a high- molecular-weight and non-crosslinked polymer additive including, but not limited to, polyethylene oxide) (PEG) in an amount ranging from 0% to about 10%, and preferably about 5% by weight.

[0141] Referring again to FIG. 1 , the electrode 100 may further an out-of-plane conductive microfiber layer 108 attached to on-textile base electrode 102 with an ionic binder layer 106. In various aspects, the microfiber layer 108 comprises a microfiber pad in which the microfibers are coated with a conductive PEDOT:PSS polymer. In some aspects, the microfiber pad is dipped in an ink comprising the conductive PEDOT:PSS polymer to coat the microfibers and form the conductive microfiber layer 108 as described in Example 1 herein. Without being limited to any particular theory, it is thought the conductive microfiber layer 108 provides for easy penetration through body hair and secure adhesion to the skin, thereby contributing to exceptional conformability, enlarged contact area, greatly reduced electrode-skin impedance and increased signal-to-noise ratio (SNR) in gel-free conditions, aswell as greatly suppressed motion artifacts even in the presence of strong motion during intense sports activities. In other additional aspects, the sensor 100 further includes at least one electrical lead 110 formed from a conductive material including at least one of a conductive metal or a conductive polymer such as the PEDOT:PSS polymer described herein.

[0142] In another aspect, to achieve waterproof E-textiles without compromising the mechanical properties of the textile and conductivity of the electrodes, an efficient one-step surface treatment can be introduced to vaporize 1 H,1 H,2H,2H-perfluorooctyltrichlorosilane (PFOTS) to form a superhydrophobic self-assembled monolayer (SAM) on the E-textile surface to enable the system to be used even in underwater conditions. A custom data recording circuit has also been designed to provide real-time signal recording, processing, and wireless communication, which in combination with the E- textile, enables on-body recording of a wide range of biopotential signals, including ECG for heart and EMG for skeleton and smooth muscles. The waterproof and perspiration-tolerant E-textiles can be made into smart garments (cycling jersey and shorts, swimsuit, etc.) to provide profound insights into athletic performance and health conditions for sport physiology applications. Moreover, clinical patient studies have also been carried out using a multi-channel E-textile system to validate the feasibility of real-time wireless monitoring of maternal ECG and uterine contractions during labor and delivery.

[0143] In various aspects, conductive polymers including, but not limited to, poly(3,4-ethylenedioxythiophene) poly(styrene sulfonate) (PEDOT:PSS)may be used to form the textile electrodes and to coat the microfiber electrode layer as described in the examples herein. In some aspects, the PEDOT:PSS conductive polymers comprise monomers with the chemical structure:

[0144] In various aspects, the 3D E-textile sensors may be waterproofed by functionalizing the sensors with a superhydrophobic monolayer formed from a superhydrophobic material including, but not limited to, 1 H,1 H,2H,2H- perfluorooctyltrichlorosilane (PFOTS). In some aspects, the PFOTS monolayer comprises the chemical structure:

[0145] The term “mmol”, as used herein, is intended to mean millimole. The term “equiv”, as used herein, is intended to mean equivalent. The term “mL”, as used herein, is intended to mean milliliter. The term “g”, as used herein, is intended to mean gram. The term “kg”, as used herein, is intended to mean kilogram. The term “pg”, as used herein, is intended to mean micrograms. Theterm “h”, as used herein, is intended to mean hour. The term “min”, as used herein, is intended to mean minute. The term “M”, as used herein, is intended to mean molar. The term "pL", as used herein, is intended to mean microliter. The term “pM”, as used herein, is intended to mean micromolar. The term “nM”, as used herein, is intended to mean nanomolar. The term “N”, as used herein, is intended to mean normal. The term “amu”, as used herein, is intended to mean atomic mass unit. The term “°C”, as used herein, is intended to mean degree Celsius. The term “wt / wt”, as used herein, is intended to mean weight / weight. The term “v / v”, as used herein, is intended to mean volume / volume. The term “MS”, as used herein, is intended to mean mass spectroscopy. The term “HPLC”, as used herein, is intended to mean high performance liquid chromatograph. The term “RT”, as used herein, is intended to mean room temperature. The term "e.g.", as used herein, is intended to mean example. The term “N / A”, as used herein, is intended to mean not tested.

[0146] The methods and algorithms of the invention may be enclosed in a controller or processor. Furthermore, methods and algorithms of the present invention, can be embodied as a computer implemented method or methods for performing such computer-implemented method or methods, and can also be embodied in the form of a tangible or non-transitory computer readable storage medium containing a computer program or other machine-readable instructions (herein “computer program”), wherein when the computer program is loaded into a computer or other processor (herein “computer”) and / or is executed by the computer, the computer becomes an apparatus for practicing the method or methods. Storage media for containing such computer program include, for example, floppy disks and diskettes, compact disk (CD)-ROMs (whether or not writeable), DVD digital disks, RAM and ROM memories, computer hard drives and back-up drives, external hard drives, “thumb” drives, and any other storage medium readable by a computer. The method or methods can also be embodied in the form of a computer program, for example, whether stored in a storage medium or transmitted over a transmission medium such as electrical conductors, fiber optics or other light conductors, or by electromagnetic radiation, wherein when the computerprogram is loaded into a computer and / or is executed by the computer, the computer becomes an apparatus for practicing the method or methods. The method or methods may be implemented on a general-purpose microprocessor or on a digital processor specifically configured to practice the process or processes. When a general-purpose microprocessor is employed, the computer program code configures the circuitry of the microprocessor to create specific logic circuit arrangements. Storage medium readable by a computer includes medium being readable by a computer per se or by another machine that reads the computer instructions for providing those instructions to a computer for controlling its operation. Such machines may include, for example, machines for reading the storage media mentioned above.

[0147] Definitions and methods described herein are provided to better define the present disclosure and to guide those of ordinary skill in the art in the practice of the present disclosure. Unless otherwise noted, terms are to be understood according to conventional usage by those of ordinary skill in the relevant art.

[0148] In some embodiments, numbers expressing quantities of ingredients, properties such as molecular weight, reaction conditions, and so forth, used to describe and claim certain embodiments of the present disclosure are to be understood as being modified in some instances by the term “about.” In some embodiments, the term “about” is used to indicate that a value includes the standard deviation of the mean for the device or method being employed to determine the value. In some embodiments, the numerical parameters set forth in the written description and attached claims are approximations that can vary depending upon the desired properties sought to be obtained by a particular embodiment. In some embodiments, the numerical parameters should be construed in light of the number of reported significant digits and by applying ordinary rounding techniques. Notwithstanding that the numerical ranges and parameters setting forth the broad scope of some embodiments of the present disclosure are approximations, the numerical values set forth in the specific examples are reported as precisely as practicable. The numerical values presented in some embodiments of the present disclosure may contain certainerrors necessarily resulting from the standard deviation found in their respective testing measurements. The recitation of ranges of values herein is merely intended to serve as a shorthand method of referring individually to each separate value falling within the range. Unless otherwise indicated herein, each individual value is incorporated into the specification as if it were individually recited herein. The recitation of discrete values is understood to include ranges between each value.

[0149] In some embodiments, the terms “a” and “an” and “the” and similar references used in the context of describing a particular embodiment (especially in the context of certain of the following claims) can be construed to cover both the singular and the plural, unless specifically noted otherwise. In some embodiments, the term “or” as used herein, including the claims, is used to mean “and / or” unless explicitly indicated to refer to alternatives only or the alternatives are mutually exclusive.

[0150] The terms “comprise,” “have” and “include” are open-ended linking verbs. Any forms or tenses of one or more of these verbs, such as “comprises,” “comprising,” “has,” “having,” “includes” and “including,” are also open-ended. For example, any method that “comprises,” “has” or “includes” one or more steps is not limited to possessing only those one or more steps and can also cover other unlisted steps. Similarly, any composition or device that “comprises,” “has” or “includes” one or more features is not limited to possessing only those one or more features and can cover other unlisted features.

[0151] All methods described herein can be performed in any suitable order unless otherwise indicated herein or otherwise clearly contradicted by context. The use of any and all examples, or exemplary language (e.g., “such as”) provided with respect to certain embodiments herein is intended merely to better illuminate the present disclosure and does not pose a limitation on the scope of the present disclosure otherwise claimed. No language in the specification should be construed as indicating any non-claimed element essential to the practice of the present disclosure.

[0152] Groupings of alternative elements or embodiments of the present disclosure disclosed herein are not to be construed as limitations. Each group member can be referred to and claimed individually or in any combination with other members of the group or other elements found herein. One or more members of a group can be included in, or deleted from, a group for reasons of convenience or patentability. When any such inclusion or deletion occurs, the specification is herein deemed to contain the group as modified thus fulfilling the written description of all Markush groups used in the appended claims.

[0153] All publications, patents, patent applications, and other references cited in this application are incorporated herein by reference in their entirety for all purposes to the same extent as if each individual publication, patent, patent application or other reference was specifically and individually indicated to be incorporated by reference in its entirety for all purposes. Citation of a reference herein shall not be construed as an admission that such is prior art to the present disclosure.

[0154] Having described the present disclosure in detail, it will be apparent that modifications, variations, and equivalent embodiments are possible without departing the scope of the present disclosure defined in the appended claims. Furthermore, it should be appreciated that all examples in the present disclosure are provided as non-limiting examples.EXAMPLES

[0155] The following non-limiting examples are provided to further illustrate the present disclosure. It should be appreciated by those of skill in the art that the techniques disclosed in the examples that follow represent approaches the inventors have found function well in the practice of the present disclosure, and thus can be considered to constitute examples of modes for its practice. However, those of skill in the art should, in light of the present disclosure, appreciate that many changes can be made in the specific embodiments that are disclosed and still obtain a like or similar result without departing from the spirit and scope of the present disclosure.EXAMPLE 1 - WATERPROOF AND GEL-FREE 3D E-TEXTILE SYSTEMS FOR EXERCISE PHYSIOLOGY AND CLINICAL MATERNAL HEALTH MONITORINGExample 1: Waterproof and cel-free 3D microfiber-on-textile electrode

[0156] To develop and validate the waterproof and gel-free 3D microfiber-on- textile electrode disclosed herein, the following experiments were conducted.

[0157] Conductive polymer poly(3,4-ethylenedioxythiophene) poly(styrene sulfonate) (PEDOT:PSS) offers solution processability and biocompatibility, and it has been extensively adopted as conductors for building soft and wearable electronics. Polyethylene oxide) (PEO), as a high-molecular-weight and non-crosslinked polymer additive, has been shown to enhance the stretchability, conductivity, and printability of PEDOT:PSS conductors (FIGS. 12B and 12C). Therefore, PEDOT:PSS / PEO composite ink has been formulated for screen-printed electrodes on textiles.

[0158] Poly(3,4-ethylenedioxythiphene)-poly (styrene sulfonate) (PEDOT:PSS, 1.3% conductive grade and 5.0% screen printing grade), polyethylene oxide) (PEO, Mv»5 000 000), Triton X-100, ethylene glycol (EG), N,N- dimethylformamide (DMF, 99.8%) were purchased from Sigma- Aldrich.1H, 7 / - / , 2 / - / , 2 / - / -Perfluorooctyltrichlorosilane (PFOTS, 97%) was purchased from thermo scientific. Silver epoxy kit (8331 D) was purchased from MG Chemicals. Silver wire (diameter ~0.004 cm) was procured from Surepure Chemetals. The nylon-spandex knitted fabric (80% nylon and 20% spandex) and polyester microfibers were purchased from FabricLA and Mayshine, respectively.TechniCloth® (TX609, blend of 55% cellulose and 45% polyester) were purchased from Texwipe. Gelled Ag / AgCI electrodes and ionic binder pads were provided by 3MTMand Hollywog.

[0159] For the electrodes printed on the textiles, the as-purchased screen printable PEDOT:PSS (5%) solution was diluted to a concentration of 3.3% using deionized water. The PEO solution was prepared by dissolving PEO in DMF with a concentration of 10 mg / mL. The mixture was then stirred at 60 °C until no precipitate could be observed. The PEDOT:PSS and PEO solutions were mixed in a weight ratio of 2:1 , followed by stirring at 80 °C for an hour to obtaina homogeneous PEDOT:PSS / PEO composite ink, denoted as PEDOT:PSS- textile ink.

[0160] For coating conductive microfibers, the as-purchased conductive aqueous PEDOT:PSS (1 .3%) was mixed with the PEO solution, Triton X-100, and EG in a weight ratio of 10:2: 1 : 1 , followed by stirring at 80 °C for an hour to produce the composite ink. The inclusion of Triton X-100 and EG additives in the ink serves multiple purposes, including tuning the rheology and wetting property of the ink for optimal dip-coating processes, enhancing ink conductivity, promoting uniform dispersion of the constituents, and improving overall electrical performance. Moreover, EG facilitates better adhesion of conductive coatings to microfibers, minimizing the risk of peel-off. The composite ink is denoted as PEDOT:PSS-microfiber ink.

[0161] On-textile electrodes were fabricated by screen-printing. Firstly, the as- purchased nylon-spandex fabric was pre-treated with oxygen plasma (60W, 2 min) to clean and improve the surface wettability. Next, a stencil mask was made by customizing the desired patterns on a polyethylene terephthalate (PET) film. A stripe pattern with dimensions of 5 mm and 14 mm in width and length, respectively, was used. Approximately 0.37 g of PEDOT:PSS-textile ink was then deposited and screen printed on to the fabric. The sample was subsequently transferred to an oven and subjected to drying and annealing at 80 °C for 1 hour. Silver epoxy and silver wire were utilized for wiring and electrical characterization.

[0162] As illustrated in FIG. 2D, the textile patch with printed electrodes was affixed to the inner bottom of a petri dish. A 10 pL droplet of PFOTS was dispensed onto a hotplate set at 90 °C. Simultaneously, the petri dish was immediately positioned on top of the droplet to create a sealed vaporization chamber. The setup is maintained for minutes to allow the PFOTS to form a uniform SAM on the textile surface (FIG. 2D, left). The sample was then flipped to treat the other side to ensure comprehensive waterproofing.

[0163] To assemble the microfiber-on-textile electrodes, the microfiber fabric was cut into pads of the desired size, followed by pre-treatment with oxygenplasma (60W, 2 min) to improve the ink absorbability. The microfiber pads were then immersed in the as-prepared PEDOT:PSS-microfiber ink to allow them to soak. Afterward, hot air was blown onto the pads using a hair dryer for predrying. The sample was thoroughly dried and annealed in an oven of 80 °C for 1 hour. Lastly, the PEDOT:PSS-soaked microfiber pad was pressed onto of the on-textile electrodes using the ionic binder pad to form the combined microfiber-on-textile electrode module.

[0164] Microstructure images were taken using an environmental scanning electron microscope (Quattro S ESEM, Thermo Fisher Scientific). The scanning electron microscopy (SEM) image (FIG. 2A) and energy dispersive X- ray spectroscopy (EDS) (FIGS. 13C and 13F) validate the uniform coating of PEDOT:PSS / PEO composite on the textile, as compared to the pristine region. To attain good elasticity and comfortability, the nylon-spandex fabric was selected to construct the wearable E- textile. Despite the inherent biaxial stretchiness of the weft knitted yarns, the stretchability and conductive pathways of the on-textile conductors still exhibit significant variations, depending on factors such as front / back sides and wale / course directions. Notably, the PEDOT:PSS / PEO conductor printed on the front side along the wale direction (FIG. 2B and FIG. 14A) exhibited superior and robust performance, with stable sheet resistance values throughout consecutive stretching tests at 5%, 15%, and 25% strain levels (FIG. 2C). To expand the applicability of the E-textile system in harsher conditions (e.g. user perspiration or underwater), a PFOTS self-assembled monolayer (SAM) was evenly vaporized on the outermost surface (FIG. 2D and FIG. 15D). The use of PFOTS SAM preserves the natural physical properties of yams and the intrinsic conductivity of the on-textile PEDOT:PSS / PEO conductor, thus overcoming a substantial challenge that conventional waterproofing treatments used for textiles (polymer coating or lamination) would completely cover the conductor and render them nonconductive. Such a facile approach provides remarkable versatility for the waterproof treatment of various textiles, including knitted (cotton), woven (polyester), nonwoven (cleanroom wipes), and even cellulose-based paper (FIGS. 16A, 16B, 16C and 16D).

[0165] To characterize contact angles, the as-prepared textile samples were placed on the stage, and a water droplet was applied to each textile surface using a pipette. Contact angle pictures were captured using a simplified microscope setup, focusing on the horizontal direction. For the long- term contact angle test, the sample was exposed to the ambient environment and examined on a daily basis. As shown in FIG. 2E, the contact angle of a water droplet on PFOTS-functionalized nylon experienced a substantial increase from 54° to 133° and the waterproof feature could last for over 7 days in ambient conditions, even after the sample was submerged completely in water for 1 hour.

[0166] To evaluate the underwater durability of the textile samples, a wiring cable was used to connect each textile sample to a multimeter and the resistance of a pristine sample was monitored. The sample was then submerged in water and subjected to various underwater operations, including bending, 180° twisting, and 360° twisting. Throughout these operations, the changes in resistance were continuously measured to monitor the resistance evolution.

[0167] E-textile samples were similarly subjected to bending and 360° twisting underwater to showcase its excellent flexibility and durability in aqueous environment, with negligible change in sheet resistance, unlike the control sample without PFOTS treatment that exhibited a gradual degradation in the electrical property (FIG. 2F). The rapid, convenient, and eco-friendly process, requiring only micron-liter PFOTS and a minute of time, dramatically broadens the adaptability of E-textiles in a wide range of scenarios, enabling proper functioning during sweating conditions or even for water sports.

[0168] Despite the advantages of the aforementioned printed on-textile electrodes, relying solely on a planar conductive film cannot eliminate the need for conductive hydrogel in biopotential recording applications due to the high electrode-skin impedance under dry conditions (FIG. 17). To eliminate the need for gel, conductive microfibers were developed as an add-on component on the base conductor. These out-of-plane fibrous conductors spread out andconform uniformly to the skin, effectively increasing the contact area and lowering the impedance.

[0169] Considering their hairy nature, different drying strategies (FIGS. 18A, 18B, 18C, 19A, 19B, 19C, 19D, 19E, 19F, 19G, and 19H) and ink formulations (FIGS. 20A, 20B, 20C, 20D, 20E, 20F, 21 A, 21 B, 21 C, AND 22) were thoroughly investigated to obtain the highly durable and reliable conductive microfibers, thereby enabling the convenient and robust biopotential measurements from dry skin. To ensure an effective bonding and electrical conduction between the on-textile conductor and the add-on conductive microfibers, multiple approaches have been studied, including stitching with PEDOT:PSS-soaked cotton threads or metal wires, and utilizing conductive binders such as silver epoxy, carbon tape, or ionic binder (FIGS. 23A, 23B, 23C, 23D, 23E, and 23F). The end-product of the 3D microfiber-on-textile electrode is schematically illustrated in FIG. 2G, which comprises microfibers soaked in the PEDOT:PSS / PEO composite ink (5 / 1 weight ratio) (FIGS. 2H, 2I, and 24) and dried with a hair dryer, securely attached to the on- textile conductive interconnect layout using an ionic binder.

[0170] To characterize electrode-skin contact impedance, two electrodes were placed on a subject’s forearm with a separation distance of 8 cm. One electrode was a gelled Ag / AgCI reference electrode, and the second electrode was an experimental electrode that was either the planar electrode without microfibers or a 3D electrode with microfibers. The measurements involved sweeping the frequency range from 10 Hz to 103Hz. The electrode-skin contact impedance was measured through an impedance analyzer (Bode 100, Omicron Lab). A multimeter was used to measure the resistance of electrodes. Equipped with the fluffy microfiber electrodes, the skin-electrode contact impedance under dry conditions decreased by nearly one order of magnitude and is even multiple folds lower than the commercial Ag / AgCI electrode, thereby making it possible to eliminate the need for gel (FIG. 2J). Such a significant reduction in electrode-skin impedance can be attributed to the increased contact area from microfibers, localized pressure, and enhanced adherence (FIGS. 25A, 25B, and 25C). All the components in FIG. 2Gcollectively form a versatile toolkit, where the microfiber pads can be conveniently attached or removed on demand, enabling the reusability of the conductive layout on textile.Example 2: Performance of the E-textile for electrophysiological signal recording

[0171] To evaluate the performance of the E-textile sensors disclosed herein for electrophysiological signal recording, the following experiments were conducted.

[0172] To verify the effectiveness of the gel-free and waterproof E-textile, we positioned a set of electrodes on forearms to capture the ECG signal under both dry and sweating conditions (FIG. 3A). To record ECG signals, freshly made electrodes were positioned on the subject’s forearms under the desired conditions, such as dry, sweating, or gelled. The electrodes were then connected to the input pins of the homemade board with a sampling rate of 500 Hz. The oscilloscope (MSO 2004B, Tektronix) displayed the real-time recording of ECG signals. The ECG amplitudes acquired by 3D microfiber-on-textile electrodes were approximately 20-fold higher in dry condition and 15-fold higher in sweating condition, as compared to the control set of solely on-textile electrodes (FIGS. 3A and 3B).

[0173] A signal-to-noise ratio (SNR) was calculated using the equation:where Asignai represents the peak-to-peak amplitude of the desired biopotential signal and Anoise represents the peak-to-peak amplitude of the background noise. The SNR achieved by the microfiber-on-textile electrodes under dry condition was over 26.5 dB, which was substantially higher than the 8.6 dB obtained by the on-textile electrodes (FIG. 3C). Remarkably, the presence of ion-rich sweat did not diminish the signal quality; instead, it enhanced the SNR by bridging the skin and electrodes through the formation of an ionic interlayer, essentially functioning similarly to the conventional gel(FIGS. 25A, 25B, 25C, 26A, 26B, and 26C). For practical daily usage, the washability of the waterproof E-textile with printed conductive on-textile electrodes is crucial for its reusability as a base platform. To simulate laundry conditions, the samples were subjected to agitation in warm water for 10 min, followed by drying with a hair dryer. The SNR remained stable after 10 washing cycles, with only small fluctuations within a range of 9~12 dB (FIGS. 3E and 3F), which validates the long- lasting waterproofing durability achieved through PFOTS treatment (FIGS. 27A and 27B). Additionally, the microfiber- on-textile electrodes were subjected to testing and storage in ambient air for a duration of over 4 months, during which no noticeable degradation in SNR was observed, demonstrating the superior long-term stability and reliability (FIGS. 3G and 3H). Furthermore, motion artifacts pose a vital challenge in the seamless integration of E- textiles with the human body, especially under gel-free conditions. Compared with the planar electrodes, the fluffy and conductive microfibers help conform to the skin robustly, functioning as a shock absorber by reducing friction and triboelectrification caused by body movement, thus effectively mitigating the motion artifacts. (FIGS. 28A, 28B, and 28C).

[0174] The E-textile is also capable of monitoring the electrical responses of muscles during nerve stimulation. EMG recording was conducted on skeleton muscle by placing newly prepared electrodes placed on the subject’s arm, with a separation distance of approximately 5 cm along the bicep. The electrodes were subsequently connected to the homemade board with a sampling rate of 500 Hz. To generate diverse EMG signals, the subject performed repetitive lifting and releasing of different weights, thereby inducing various levels of muscle contractions. The oscilloscope provided a real-time display of the recorded EMG signals. SNR of the EMG signals was determined as described above.

[0175] The microfiber-on-textile electrodes, spaced 5 cm apart, were positioned on the arm to collect bicep-EMG signals (FIG. 3I). As the weight being lifted was increased, the EMG amplitudes increased accordingly (FIGS. 3J and 3K) and the SNRs under dry and sweating conditions were comparable (FIG. 3L), further indicating the gel-free and waterproof features of our E-textile. Notably, our microfiber-on-textile electrodes consistently outperformed the commercialAg / AgCI electrodes in terms of sensitivity, signal amplitude and SNR across all tests (FIGS. 29A and 29B).

[0176] Considering factors such as ECG amplitude and SNR, the microfiber-on- textile electrodes exhibited enhanced performance compared to the planar on- textile electrodes in all conditions. Comparing the planar on-textile electrodes used in dry and sweating conditions, the significant increase in SNR observed under sweating conditions helped confirm the poor electrode-skin contact interface in dry conditions, as well as the sweat tolerance of PFOTS-treated E- textiles. On the other hand, unlike the planar on-textile electrode, the 3D microfiber-on-textile electrodes demonstrated comparable signal amplitudes and high SNR in all conditions, validating their effectiveness and durability for potential all-day wearing usage in gel-free conditions.

[0177] When using the commercial Ag / AgCI electrodes for biopotential recording, the electrodes are typically required to be firmly fixed at the desired location using 3M tapes and are thus usually capable of producing clean and stable signals when the subject is stationary. However, when subjected to body movements, the intense friction at the electrode-skin interface could result in significant motion artifacts that might interfere with the recorded signals. For the planar on-textile electrodes, poor contact with dry skin and high electrodeskin impedance led to noticeable motion artifacts caused by the subject's routine breathing activity even when the subject is stationary. In contrast, the 3D microfiber-on-textile electrodes exhibited good tolerance to motion artifacts and were able to record stable signals during both static and active movements (FIGS. 28A, 28B, and 28C). Moreover, despite the presence of small motion artifacts when the subject was in motion, distinguishable ECG characteristic peaks were clearly captured.Example 3: Real-time ECG & EMG monitoring during intense cycling exercise

[0178] To evaluate the performance of the E-textile sensors disclosed herein for electrophysiological signal recording during cycling, the following experiments were conducted.

[0179] Sports physiology relies on extensive real-time data collection (e.g. heartrate, power output, VO2 Max, etc.) to guide athletes’ training for enhanced performance. While real-time electrophysiological signal recording is invaluable, it presents significant challenges, particularly during intense exercise, where motion artifacts and user perspiration can pose issues. The 3D microfiber-on-textile electrodes are ideally suited for such applications. A single-channel wireless data recording system on miniaturized printed circuit board (mini-PCB) has been designed and manufactured to connect with the E- textile, thus enabling a compact and wearable system for on-demand placement, real-time signal collection and processing, and wireless communication with a user interface through a radiofrequency (RF) module (FIG. 4A). The detailed circuit diagram of the analog frontend and the cutoff frequency calculation are illustrated in FIG. 30.

[0180] To demonstrate the practical application of the E-textiles for exercise physiology, a cyclist wore a cycling jersey and cycling shorts equipped with microfiber-on-textile electrodes while cycling on a stationary bike. The cycling jersey was equipped with paired microfiber-on-textile electrodes symmetrically positioned on the inner side of the jersey facing the lower chest. The cycling shorts were equipped with paired microfiber-on-textile electrodes positioned vertically and facing a single quadricep with a separation distance of approximately 5 cm.

[0181] For ECG monitoring during the cycling training, one set of on-textile electrodes was printed on the inner side of the cycling jersey, symmetrically positioned towards the chest of the subject, and the electrodes were connected to the single-channel mini-PCB. When the cyclist was ready to begin the training session on the stationary bike (Peloton Bike), the microfiber pads were attached onto the printed electrodes to achieve the microfiber-on-textile electrodes for gel-less recording. The recording was be initiated or stopped by turning the power supply on or off using a switch on the coin cell battery holder. The real-time ECG recording during the training sessions was displayed on the tablet through a plug-in receiver (FIG. 4A). After completing the entire training session, the disposable microfiber pads were peeled off and discarded.

[0182] For EMG monitoring during the cycling training, one set of on-textile electrodes was printed on the inner side of the shorts facing the thigh, spaced 5 cm apart along the single quadricep, which were connected to the singlechannel mini-PCB. Following a similar procedure of attaching microfiber pads and switching on the power supply, the real-time EMG recording was displayed on the tablet during the cycling training (FIG. 4A).

[0183] For the cycling experiments, the E-textile system was programmed with a sampling rate of 500 Hz to ensure accurate signal capture and fidelity. In the case of single-channel applications, specific signal processing techniques were employed for ECG and muscle EMG signals.

[0184] For ECG signal processing, a bandpass filter with a frequency range of 1 Hz to 50 Hz was applied to filter out noise and unwanted frequency components. Subsequently, a Wavelet analysis was performed to extract the R peaks of the recorded ECG signals. By measuring the time intervals of ten consecutive R peaks detected, the heart rate was calculated. To provide a continuous heart rate measurement, a moving step of one was applied, and the results are plotted in FIG. 4B (bottom graph). Regarding the muscle EMG signals, they were extracted using a bandpass filter with a frequency range of 50 Hz to 240 Hz. To derive an instantaneous cycling cadence, the time intervals between fifteen EMG signal peaks were measured. A continuous cycling cadence was obtained by employing a sliding window with a step size of one, and the results are shown in FIG. 4C (middle graph). The EMG signal envelopes were processed using a sliding window of 150 samples. The EMG signal amplitude is detected as the envelope peaks and an amplitude plot is shown in FIG. 4C (bottom graph).

[0185] Continuous and real-time monitoring of ECG signals was conducted throughout a 14-minute training session, comprising a 4.5-minute warm-up phase, a 5-minute climbing phase, and a 4.5-minute sprinting phase (FIG. 4B), as well as heart rate (HR) analysis was performed to monitor the real-time HR for zone-based training.

[0186] During the initial stage, the subject began in a relaxed state of pedalingfor warming up and the recorded ECG signals exhibited exceptional stability (FIG. 4B), where the HR averaged at around 120 beats per minute (bpm). As the climbing phase began, the subject exerted greater efforts and lowered the upper body to overcome the increased pedaling resistance, resulting in a gradual rise in HR to 160 bpm. The motion artifacts remained negligible in this phase with only occasional voltage spikes that had no discernible influence on the signal quality. During the sprinting phase, the subject had to move his body left-and-right vigorously and the presence of motion artifacts became more noticeable. Nevertheless, even under such extreme condition, the unprocessed raw ECG signal still preserves all distinct ECG features with clear QRS complex and T-wave (FIG. 4B). Moreover, despite the subject’s intense perspiration throughout the climbing and sprinting phases (FIG. 4B), the waterproof durability of the E-textile enabled uninterrupted recording.

[0187] EMG is another crucial electrophysiological signal that athletes monitor to assess muscle output and track training performance. During the experiment, the cyclist wore a E-textile-based cycling short with microfiber-on- textile electrodes positioned 5 cm apart facing the thigh to monitor the electrical activities of the quadriceps muscle. The thigh-EMG signals were monitored in real-time during a 6-minute session, comprising four phases: warm-up, climbing, sprinting, and cool-down. It is worth noting that performing on-body real-time EMG recording, particularly in gel-free conditions, poses extreme challenges due to intense frictions and deformations between the skin and the electrodes caused by vigorous body movements and localized muscle contractions. The raw thigh-EMG signals collected by our E-textile and the zoomed-in views (FIG. 4C) clearly indicated the evolution of muscle contraction activities during different exercise phases. Meanwhile, additional data analysis was conducted to extract valuable insights, allowing the cyclist to make timely and professional assessments of exercise performance and muscle activity. Cadence is a vital metric in cycling training used to track the pedaling rate to guide the cyclists in maintaining effective power output and minimizing the risk of overuse injury. To derive cadence information from the raw EMG signal, the following steps were employed: applying the bandpass filtering (1~50 Hz) toisolate the target frequency range, utilizing the maximal overlap discrete wavelet transform (MODWT) algorithm to do segmentation, identifying the peaks corresponding to muscle contractions, and calculating the interval between consecutive peaks as the duration per pedal revolution. The cadence extracted by the E-textile precisely aligned with the data recorded by the commercial sensing module mounted on the stationary bike’s crank arm (FIG. 4C, center graph). Besides cadence, the EMG amplitude analysis provides information about the muscle activation and exertion during the training. The extracted EMG amplitude exhibited a proportional correlation with the digital resistance setting on the bike (FIG. 4C, bottom graph). During the warm-up phase of low resistance pedaling, the EMG amplitude remains fairly low and stable with clear plateau patterns that match the bike’s resistance setting, indicating sustained and consistent muscle output for generating a constant force. When dialing up the resistance to enter the climbing phase, the subtle delay followed by an abrupt rise in the EMG amplitude suggests the presence of hysteresis in muscle activation and neuromuscular coordination. When the subject was struggling in the sprinting phase, the fluctuation and instability in the EMG amplitude could be attributed to intensified muscle contraction and vigorous body swinging required to overcome the over-loaded resistance. This unstable pattern persists until reaching volitional fatigue, followed by the transition into the cool- down phase with a drop in EMG amplitudes and subsequent stability.Example 4: Real-time ECG monitoring during swimming exercise

[0188] To evaluate the performance of the E-textile sensors disclosed herein for electrophysiological signal recording during swimming, the following experiments were conducted.

[0189] Despite commercial products such as swim watches, chest straps, and arm bands have emerged for monitoring heart rate (HR) during water sports, they are limited to recording HR at discrete time intervals. Achieving on-body real-time and continuous recording of raw electrophysiological signals in an underwater environment remains an unsolved challenge. The substantial interference caused by water renders the use of electrodes underwaterimpractical, further complicating the accurate biopotential recording.

[0190] In this experiment, the disclosed waterproof and gel-free E-textile system was integrated into a swimsuit, and real-time and continuous monitoring of the swimmer’s ECG was conducted. A swimmer wore a swimsuit provided with paired microfiber-on-textile electrodes positioned horizontally at the chest-level, facing the swimmer’s back.

[0191] For the swimming experiments, the E-textile system was programmed with a sampling rate of 500 Hz to ensure accurate signal capture and fidelity. The specific signal processing techniques for ECG and muscle EMG signals as described in Example 3 above were used in the swimming experiments.

[0192] One set of electrodes was printed on the inner side of the swimsuit, symmetrically positioned to face the subject’s back at the chest level. The electrodes were connected to the mini-PCB, which was securely packaged in a waterproof pocket around the waist of the subject. The subject put on the swimsuit in the locker room and then switched on the system for recording. After powering up the recording system, the swimmer took a brief shower and walked to the pool to begin the swimming training. The wireless data recording followed a similar procedure as the cycling test as described in Example 3, with a tablet placed poolside to display real-time ECG recordings of the swimmer.

[0193] An eleven-minute recording session was conducted that included various activities, including walking from the locker room to the pool (FIG. 5A), stepping into the pool (FIG. 5B), floating and diving (FIG. 5C), swimming (FIG. 5D), exiting the pool (FIG. 5E), and walking back to the locker room (FIG. 5F). During the initial 3.5 minutes, after putting on the E-textile-equipped swimsuit and a brief shower to moisturize the skin, the swimmer proceeded to walk towards the pool, during which a nearly motion-artifact-free ECG recording could be observed (FIG. 5A). When stepping into the pool, significant motion was introduced during the process, which coupled with water gradually permeating through the swimsuit, causes observable interference with the sensing electrodes facing the swimmer’s back. The interaction with water and friction at the electrode-skin interface resulted in motion artifacts and noticeablebut small voltage spikes. Nevertheless, upon closer examination of the raw signals captured, the characteristic ECG features such as the QRS complex remained distinct even during the noisiest phase of the process (FIG. 5B). During floating and diving practices to acclimate to the water environment and warm up, the sensing module recorded consistently without disruption even when the swimmer was completed submerged underwater (FIG. 5C). Afterwards, the participant performed normal swimming activities using breaststroke style along the lane for several rounds, and the recorded signal quality remained unaffected by the intense body motions and water interference (FIG. 5D). As the swimmer exited the pool, minor motion artifacts could once again be observed in the recording, but the ECG signals immediately recovered and stabilized during the walk back to the locker room (FIGS. 5E and 5F). Overall, the successful implementation of E-textile for underwater electrophysiological signal recording showcases a remarkable advancement in technology for water sports physiology, providing a promising tool for elevating training effectiveness.Example 5 Real-time clinical maternal health monitoring during labor

[0194] To evaluate the performance of the E-textile sensors disclosed herein for electrophysiological signal recording during maternal labor, the following experiments were conducted.

[0195] In obstetrics, uterine contractions are usually recorded clinically using pressure-sensitive tocodynamometer (TOCO) or intrauterine pressure catheter (IUPC). However, these “gold- standard” instruments sometimes have limited signal accuracies due to incorrect instrument placement and causes significant patient discomfort due to the tight bandages used for TOCO or the invasiveness from IUPC. The uterine EMG signal plays an essential role in assessing and interpreting the uterine contraction activities, providing valuable aid in diagnosing labor dysfunction and predicting preterm labor. However, real-time acquisition of uterine EMG signals can be challenging due to the dynamic nature of uterine contractions and the complexity of achieving acceptable SNR due to the low EMG amplitude produced by smooth muscles, the deep- lying location of the uterus, the interference by electrophysiologicalsignals from neighboring organs and heart, and the maternal body movement. Recently, a new high-resolution and noninvasive electromyometrial imaging technology has been developed, which uses an array of unipolar electrodes to measure uterine electrograms from the patient's abdomen surface to generate three-dimensional maps of uterine electrical activities during the labor contraction.

[0196] A total of four pregnant subjects participated and completed the entire study. These subjects represented a diverse range of characteristics, including variations in age (19~28 years old), delivery history (nulliparous or multiparous), labor type (induction or spontaneous), and body conditions (body mass index 26.21 ~39.88). Detailed descriptions of the four subjects of these experiments are provided in Table 1 below.

[0197] TechniCloth® was selected as the textile-based substrate for these experiments, which offered the lowest level of releasable ions and non-volatile contaminants, contained no chemical binders and tolerated most standard solvents. A mesh-type mask was customized (made by Tokyo Process Service Co., LTD) for the screen printing of electrode array layout, which would be used to manufacture multi-channel E-patch for clinical monitoring of maternal health (FIG. 34). On average, 2 grams of PEDOT:PSS-screen-printing ink would be enough for manufacturing one single patch. The annealed patch was then wired with ribbon cable using silver epoxy as binder. A medical-grade double side tape (3MTM) was used for both adhesion to human skin and encapsulation of the sensor patch with circular openings to expose only the active sensing regions.

[0198] To use the E- textile system in uterine contraction monitoring, the singlechannel recording module described in Examples 3 and 4 above was upgraded to a multi-channel version with additional multiplexing functionality at the frontend (FIG. 6A). This instrumentation integrates a separate daughter-PCB beneath the mother-PCB to incorporate a differential analog multiplexer and active shielding circuits, thereby enabling simultaneous recording from up to 16 channels (FIG. 31 ). An E-textile patch with a printed electrode array layout(FIG. 6B) connected to the stacked PCBs was attached to the maternal abdomen during the experiment. This configuration also offers wireless transmission of signals to a receiving device located outside the labor & delivery room, enabling real-time multi-channel signal recording, monitoring, and analysis.

[0199] The clinical studies were conducted in labor & delivery rooms at a treatment facility with the assistance of doctors, nurses, and clinical assistants. The experiments were performed when the subjects were in active labor, with a cervical dilation ranging from approximately 3.5 to 4.5 cm. Typically, multiple devices, such as TOCO monitors and fetal monitors had already been placed on the subject’s abdomen to monitor uterine contractions and fetal HR, providing real-time monitoring to guide the decision-making in accordance with clinical practices. However, in some cases when the noninvasive TOCO system failed to effectively capture the contraction signals, an invasive intrauterine pressure catheter (IUPC) was applied to the subject. To ensure the safety of both the mother and fetus, the clinical patient studies commenced only after confirming regular maternal uterine contractions and fetal health.

[0200] After confirming the clear uterine contraction signals from either the TOCO or IUPC systems, an E-textile patch (FIG. 6B)was connected to the input pins of the multi-channel data recording system, and the release liner of the double-sided medical tape was removed to render the sensor patch ready for use. The sensor patch was securely attached to the subject’s abdomen in a position adjacent to the TOCO monitor to minimize spatial delay of uterine contraction signals. A commercial high-resolution BioSemi system, with active Ag / AgCI electrodes (BioSemi B.V, Amsterdam, The Netherland), was additionally applied to occupy the remaining available space on the abdomen. The clocks of both the E- textile and BioSemi systems were synchronized, and recording was initiated simultaneously. A tablet was placed outside the labor & delivery room to display the real-time signals captured by the wireless E-textile system, while a laptop was utilized to display the recordings from the BioSemi system using a wired optical fiber. Typically, four separate 15-minute recording sessions were conducted by entering the labor & delivery room and obtainingpermission from the subject to proceed. Once all the sessions were completed, the systems were switched off, and all the patches were carefully removed from the subject's abdomen. Each used E-patch was discarded directly, while the used BioSemi patches underwent a sequence of cleaning, drying, and tape replacement.

[0201] During the clinical patient study, multiple instruments, including an 8- channel E-textile system, the TOCO system, and a commercial biopotential measurement system with active Ag / AgCI electrodes (BioSemi), were implemented together on pregnant subjects to cross-validate the effectiveness and quality of recordings. In addition to uterine EMG, maternal ECG could also be captured simultaneously from the same electrodes, revealing the coupling of multiple biopotential signals originating from various organs and tissues embedded in the raw recording. Therefore, a dedicated protocol for data interpretation has been developed to enable multimodal monitoring of maternal health. As illustrated in FIG. 8A, the wirelessly received signals were initially demultiplexed into separate channels, followed by median filtering, crosstalk compensation, bandpass filtering, and motion-artifact cancelation to enhance the signal quality (FIG. 6A). Fast Fourier Transform (FFT) analysis was performed to extract feature components in frequency domain (FIG. 8B), ultimately contributing to the identification of dominant frequency peaks and distributions associated with maternal ECG (1~3 Hz) and uterine EMG (0.3~1 Hz). Following this protocol, the real-time ECG recorded throughout the entire 21- minute clinical session is presented in FIG. 6C (top graph), revealing high- quality signals with clear ECG features in the zoomed-in windows. The realtime HR analysis, averaging at 87 bpm, reflects regular cardiac cycles and rhythm of the subject and aligns well with the dominant frequency (1 ,45~1 .5 Hz) observed in the time-frequency analysis (FIG. 6C).

[0202] Regarding the uterine EMG monitoring, eight uterine contractions were confirmed using the TOCO, along with some irregular spikes likely caused by maternal movements (FIG. 6D). The raw EMGs recorded by the E-textile and BioSemi closely resemble each other in terms of amplitude and SNR.Moreover, the root-mean-square (RMS) waveform derived from the E-textilesignal correlate excellently with the TOCO waveform. To validate the reproducibility of E-textile in clinical settings, it was tested in a total of four clinical patient studies by involving multiple pregnant subjects (Subjects #1-4). The E-textile successfully accomplished all patient studies, yielding high-quality multimodal recordings. In contrast, the commercial BioSemi using rigid electrodes missed some feature contractions for Subject #2 (FIG. 9B) and failed to record due to the patch delamination caused by Subject #3’s vigorous movements (FIG. 10B). The TOCO also failed to record due to the displacement induced by Subject #4’s motion (FIG. 11 B). Notably, these four patients represented a diverse range of characteristics, including variations in age (19~28 years old), delivery history (nulliparous or multiparous), labor type (induction or spontaneous), and body conditions (body mass index), providing compelling evidence to evaluate the effectiveness and robustness of the E- textile, as summarized in Tables 1 and 2.Table 1: Subject information for clinical studies.Table 2: Performance evaluation of the multiple instrumentation used for uterine-EMG monitoring across all subjects during clinical studies.

[0203] Moreover, the E-textile’s ready-to-use and disposable attribute eliminate tedious workload associated with the electrode preparation and post-cleaning needed for the BioSemi in clinical studies (FIGS. 33A and 33B). This simplifies the overall usage procedure and could potentially allow users to apply the E- textile themselves for in-home monitoring applications. More significantly, when compared to the state-of-the-art uterine contraction monitoring equipment that costs over tens of thousands of US dollars, our E-textile system stands out as a highly affordable alternative. The reusable wireless recording hardware costs around $93, while the disposable patch costs less than $5. The affordability of this low-cost E-textile system paves the way for widespread adoption of the technology for both clinical and in-home maternal health monitoring and labor management practices, especially in low-resource settings.

[0204] Typically, the uterine contractions commence in proximity to the expected delivery date and may persist for extended hours as childbirth approaches. To ensure the long-term monitoring feasibility of the E-textile system, an hour-long (approximately 67 minutes, as restricted by the IRB approval) clinical study was conducted. A total of 22 contractions were identified throughout the recording period (FIG. 7A). The segment encompassing the contraction #20 was specifically magnified to examine the multi-channel uterine EMGs acquired by the E-textile (FIG. 7B). The temporal transitions of electrical activation / deactivation were tentatively inferred from EMG-RMS waveform (denoted by the blue-red-blue step line). By rendering the EMG recorded by each individual electrode at discrete time points, the dynamic potential maps on body surface throughout the contraction #20 was generated.Sequential potential maps at indicated time windows (t1 to t4) revealed the temporal and spatial evolution during the uterine electrogram burst (UEB). The feasibility of mapping uterine contractions by the multi-channel E- textile was further substantiated by incorporating with the BioSemi to visualize electrical propagation patterns. Periodic positive and negative electrogram potentials during the UEB could be observed from spatial-temporal potential maps, illustrating evident radial propagation originating from approximate TOCO- placed area. Notably, the same spatial resolution (~2.5 cm) has been achieved by E-textile system, which could be readily increased by adjusting the electrode layout.

[0205] Analysis of the extracted contraction intervals from the E-textile indicated a gradual decrease from 6 minutes to less than 2 minutes, reflecting the trend of contractions becoming more frequent (FIG. 7C). Frequent contractions, especially when they become stronger and closer, typically signify the progression of labor, aiding in the descent of the baby through the birth canal. The precision of interval analysis was cross validated with the TOCO, yielding an accuracy of over 95% (FIGS. 7C, 32A, and 32B).Example 6: Drying methods of conductive microfibers

[0206] To evaluate the impact of washing and drying the E-textile sensors disclosed herein on electrophysiological signal recording performance, the following experiments were conducted.

[0207] Given that the conductive microfibers will function as electrodes in direct contact with human skin, their robustness and durability are significant metrics to be considered. To achieve a PEDOT:PSS coating on the microfibers in a fast and cost-effective manner, a dip-coating technique was developed. However, a challenge arises when it comes to drying and annealing the dipped microfibers. Two approaches, drying in oven (FIG. 18A) or by hair dryer (FIG. 18B), have been investigated and compared. The sample dried using a hair dryer exhibited better durability as it was able to withstand more washing cycles (FIG. 18B) and delivered a better SNR (FIG. 18C). The improved performance associated with the hair dryer approach can be attributed to thefollowing aspects:

[0208] The hair dryer allows for gentle and controlled drying. PEDOT:PSS coatings are sensitive to high temperatures, and excessive heat can cause degradation or delamination. By utilizing warm air from the hair dryer, this gentle drying process minimizes the risk of damaging the PEDOT:PSS coating, ensuring its integrity and robustness.

[0209] The airflow generated by a hair dryer aids in achieving a fluffy drying effect. Microfibers coated with PEDOT:PSS tend to clump together when wet. The airflow from a hair dryer helps to separate and fluff up the fibers, restoring their original texture. In contrast, oven drying lacks the necessary airflow, resulting in the formation of bulk and fragile PEDOT:PSS segments that are prone to breaking and delamination, especially after washing (FIGS. 19A, 19B, 19C, 19D, 19E, 19F, and 19G).

[0210] Hair dryers provide a quicker drying process, minimizing the potential risks associated with prolonged high temperature drying and preserving the PEDOT:PSS coating's robustness. Extended exposure to high temperatures during oven drying may increase the likelihood of thermal degradation that could compromise the coating's quality and the elasticity of the microfibers.Example 7: Formulation of PEDOT.PSS / PEO ink for conductive microfibers

[0211] To evaluate the effects of different conductive ink formulations on E- textile sensor performance, the following experiments were conducted.

[0212] The objective of introducing conductive microfibers is to leverage the advantages of fluffy and fibrous conductors to enable gel-free electrophysiological monitoring on dry skin. As discussed in Supplementary Note 1 , pure PEDOT:PSS coatings are brittle and lack mechanical durability, rendering them unable to withstand the significant deformations induced by the soft and flexible microfibers. Besides, the poor adhesion of pure PEDOT:PSS can lead to delamination or peeling from the microfibers. In addition, the poor adhesive nature can weaken the conformability of PEDOT:PSS-coated microfibers to human skin, thereby increasing the electrode-skin contactimpedance and reducing the biopotential recording quality, especially in dry conditions (FIGS. 20A and 20D). To address these challenges, an approach of blending PEG into the PEDOT:PSS ink has been developed to optimize the composite ink for dip coating microfibers. Two mixing ratios, 10 / 1 and 5 / 1 weight ratios of PEDOT:PSS / PEO composite ink, were investigated. Notably, the electrode-skin contact impedances obtained from PEDOT:PSS / PEO- coated microfibers in dry conditions (Supplementary FIGS. 20B and 20C) were significantly lower compared to pure PEDOT:PSS-coated microfibers (FIG. 20A). Moreover, the quality of recorded ECG signals, in terms of both raw signal amplitude and SNR, was enhanced under all conditions (FIGS. 20D, 20E, and 20F). Further characterization was conducted to assess the durability variation of different blending ratios of PEO additive in order to determine the optimal ink formulation. As shown in FIG. 21 A, microfibers coated with 10 / 1 PEDOT:PSS / PEO composite ink almost failed to produce clear EMG signals after four washing cycles, whereas microfibers coated with 5 / 1 PEDOT:PSS / PEO ink remained functional even after 12 washing cycles (FIG. 21 B), demonstrating their superior robustness and washability. Besides, the 5 / 1 sample exhibited higher SNR during the initial washing cycles, indicating better contact and lower electrode-skin impedance.

[0213] SEM images were captured to investigate the underlying reasons for this durability variation. Microstructure images were taken using an environmental scanning electron microscope (Quattro S ESEM, Thermo Fisher Scientific). As shown in FIG. 22, both 10 / 1 and 5 / 1 sample showed uniform polymer coating on individual fibers, with cross-linking junctions forming a percolated conductor network. After 4 washing cycles, significant coating delamination and lack of interconnections were observed in the 10 / 1 samples, while the fibers from the 5 / 1 sample remained well cross-linked. Even after 10 washing cycles, effective bonding among microfibers was still evident, confirming the improved durability with more PEO blended into the PEDOT:PSS. Therefore, the 5 / 1 PEDOT:PSS / PEO ink formulation was chosen for subsequent studies.Example 8: Signal processing and data analysis

[0214] To develop and validate single-channel and multiplexed signal acquisition and analysis using the E-textile sensors disclosed herein, the following experiments were conducted.

[0215] The raw signals of interest from the E-textile sensors underwent signal processing through a bandpass filtering network, which included a low-pass filter and a high-pass filter (FIGS. 30 and 31 ).

[0216] For the single-channel system, the frequency filtering band was set to [fHP_single, fLP_single], 3S defined below:

[0217] For the multiple-channel system, the frequency filtering band was set to [fnp_muiti, fLP_muiti], as defined below:

[0218] After the bandpass filtering, the signals were amplified using a high- performance instrumentation amplifier (Texas Instrument INA819). To ensure optimal amplification, the amplifier gain was set to 60dB for single-channel applications and 40dB for multi-channel applications. For the in-lab experiments, raw data were directly exported from the oscilloscope without undergoing additional post-processing, displaying signals in the Volt scale with a 60dB gain from INA. For the exercise tests and clinical studies, raw data were acquired wirelessly and subsequently subjected to post- processing, presenting signals in the milli-Volt scale after dividing the data by the 60dB gain. To facilitate digitization and further processing, a high-resolution 16-bit analog-to-digital converter (ADC, Texas Instrument ADS8689) was employed. The amplified signals were digitized by the ADC and transmitted to a microcontroller (MCU, Texas Instrument CC1310) for subsequent analysis and manipulation. The communication between the ADC and the MCU wasfacilitated through the Serial Peripheral Interface (SPI) protocol, ensuring reliable and efficient data transmission. The MCU played a crucial role in packaging the received data and wirelessly transmitting it to a remote receiver. The remote receiver then facilitated the transfer of the data to a personal computer (PC) or tablet using the universal asynchronous receiver / transmitter (LIART) communication protocol, enabling seamless data integration and analysis. To mitigate the effects of parasitic capacitance and ensure accurate signal acquisition, an active shield circuitry was incorporated. This circuitry effectively enclosed the input electrodes and canceled out the unwanted parasitic capacitance, as well as minimizing interference from the signal transmission lines. The active shield buffers were implemented using operational amplifiers (Texas Instrument OPA130), known for their high- performance characteristics and reliable operation.3.1 Single-channel data for exercise physiology experiments.

[0219] In the case of multi-channel applications, the wirelessly received data comprises all sampled channels. To ensure signal integrity, each channel is programmed with a sampling rate of 200 Hz. To extract the desired ECG and uterine contraction signals from the multi-channel data, a specific signal processing pipeline is employed, shown illustrated at FIG. 8A.

[0220] Referring to FIG. 8A, the data raw multiplexed data is received at 802 and undergoes a demultiplexing process to separate the individual channels at 804, and filtering applied to each of the separated individual channels at 806. A median filter is applied to mitigate any noise artifacts that arises from wireless couplings.

[0221] For the ECG signal extraction, a bandpass filter is designed to pass frequencies within the range of 0.5 Hz to 50 Hz. This filtering process effectively isolates the ECG components of interest from the multi-channel data. The resulting ECG signal is illustrated in FIG. 6C. To analyze the recorded ECG signals, the heart rate is extracted by detecting the time intervals between 20 consecutive R peaks in the ECG waveform. Continuous heart rates are calculated with a step size of 5 peaks and the data is presentedin FIG. 6C (middle graph). Additionally, the power spectrum of the recorded ECG signals is plotted in FIG. 6C (lower graph) to visualize the signal frequency characteristics.

[0222] For uterine contraction EMG signals, a bandpass filter with a frequency range of 0.3 Hz to 1 Hz was applied. This targeted filtering approach isolates the specific frequency band associated with uterine contractions. The root- mean-square (RMS) of the filtered signals was calculated, and unwanted motion artifacts were removed based on their higher magnitude compared to uterine contractions.

[0223] To ensure reliable and stable power delivery, premade voltage regulating circuits were employed for both the single-channel and multi- channel systems. Specifically, a voltage regulating circuit (Adafruit 4654) was utilized for the 5V supply, while another circuit (Adafruit 2795) was employed for the 3.3V supply. These voltage regulating circuits provided robust power regulation, ensuring consistent and clean power to the entire system. The single-channel system was powered by a 3.7-volt coin cell battery. The multi- channel system was powered by 4 AA batteries, whose power consumption was approximately 140 mW.

[0224] By incorporating these signals processing techniques, including demultiplexing, median filtering, and bandpass filtering, individual channels were effectively separated from the multiplexed signals, noise interference caused by wireless couplings was reduced, and the subject’s ECG and uterine contraction EMG signals were isolated for further analysis.Example 9 PEDOT.PSS / PEO blend for on-textile electrodes

[0225] To evaluate the effect of the proportion of PEG in a PEDOT:PSS / PEO blend used to fabricate the E-textile sensors disclosed herein, the following experiments were conducted.

[0226] PEDOT:PSS is well-known for its conductivity, transparency, and biocompatibility and has been extensively studied for building conductors used in biomedical-related applications. However, it is worth noting that PEDOT:PSS does have certain limitations in specific areas. For instance, the lowmechanical strength of PEDOT:PSS thin films or coatings makes them susceptible to cracking when subjected to deformation. To overcome these potential drawbacks, PEO has been investigated as an effective additive to improve the performance of PEDOT:PSS ink as reported in our prior publications2,3. It should be noted that the solvent used to dissolve PEO is polar solvent DMF, which could also be regarded as an additive that contributes to phase separation between the PEDOT and PSS grains. PSS- is negatively charged insulator which could help stabilize the positively charged conductor PEDOT+, but at the cost of hindering the charge transport pathways. By employing a polar solvent, the microstructure of the PEDOT:PSS film becomes more percolated, promoting charge hopping and thereby enhancing the film's conductivity. The commercial PEDOT: PSS ink provided an acidic environment, where the ether groups in PEO underwent protonation and cleaving effects, resulting in the formation of hydroxyl groups (Step 1-2). These hydroxyl groups readily react with the sulfonic acid groups in PSS, forming sulfonic acid esters (Step 3). Consequently, the PEO additive can be successfully cross-linked to PSS, effectively softening the PEDOT: PSS and yielding a more flexible and durable composite conductors4 (Supplementary FIG. 1 B).

[0227] Stretchability tests were conducted using a syringe pump (KDS Legato 110), whole motions could be precisely programmed. The stretchability tests of planar on-textile electrodes printed on the back side of textile along the wale direction, using as-purchased pure PEDOT:PSS ink (FIG. 12C) and PEDOT: PSS / PEO composite ink (FIG. 14C), further validate the effectiveness of the PEO additive in enhancing the durability and robustness of the electrodes.Example 10: Microstructure Analysis

[0228] To evaluate the microstructural characteristics of PEDOT: PSS-coated textiles used in the construction of the E-textile sensors disclosed herein, the following experiments were conducted.

[0229] The EDS elemental analysis conducted on the red-circled area of the threads in the pristine textile revealed no presence of sulfur on bare fibers (FIGS. 13A, 13B, and 13C). In contrast, uniform coating layers could be observed in the PEDOT:PSS-coated textile sample (FIG. 13D), and EDS elemental analysis confirms the presence and even distribution of sulfur on the microfibers (FIGS. 13E and 13F).Example 11: Current-flow models and stretchabilitv characterization

[0230] To evaluate the performance of the E-textile sensors disclosed herein under stretching conditions, the following experiments were conducted.

[0231] Current flow models were developed and used to evaluate the performance of the E-textile sensors under various stretching conditions.

[0232] The stretchability tests were conducted on planar on-textile electrodes, involving repeated stretches to 5%, 15%, and 25% strain levels, with a 5% prestretch. The current flow model was divided into the front and back sides, with each side comprising the wale and course directions. This model can be described as a linear thread with interconnecting junctions5,6. The textile was stretched along the gray arrow, and the sheet resistances were measured in the direction of current flow. Junction 1 (J1 ) and Junction 2 (J2) represented relatively good and poor junctions, respectively, based on the knitting structure.

[0233] FIG. 14A shows the most robust performance obtained by printing the electrode along the wale direction on the front side of the textile. This orientation consists of the shortest current-flow path (L = 2.74 mm) and the fewest number of poor junctions (2*J2) in the visualized area. Specifically, the thread skipped one course and connects to the next course through J2, located on the back side of the textile. Interestingly, stretching along wale direction tightened junction J2, resulting in cracks but maintaining a strong connection. Even though the threads are connected through relatively poor junctions, stretching force does not seem to significantly affect their electrical properties.

[0234] FIG. 14B (front-course) and FIG. 14C (back-wale) demonstrate similar stable sheet resistance results during stretching. The front-coursecircumstance consists of a current-flow path of L = 3.61 mm and eight good junctions (8XJ1 ) in the visualized area, while the back-wale circumstance consists of a current-flow patch of L = 2.79 mm and nine good junctions (9XJ1 ). In these orientations, the current flowed only through good junctions, resulting in relatively robust resistance. Although the front-course path was 1 mm longer than the back-wale path after stretching, the length before reaching the junction (!_' = 259.3 pm) was only half as long as that of the back-course (!_' = 496.8 pm). As a result, the current could flow with less disturbance along the frontcourse orientation.

[0235] FIG. 14D shows that the back-course orientation exhibited the most unstable resistance when stretched. This can be clearly observed in the optical microscope image and explained by the current flow model. Stretching completely disconnects the whales and created additional path length and poor junctions. Ultimately, the back-course orientation consisted of a path length of 3.46 mm, four good junctions (4XJ1 ) and nine poor junctions (9xJ2). Furthermore, the location of the new J2 continuously changes with stretching. Consequently, during the test, additional poor junctions were repeatedly disconnected and formed, leading to significant variations in resistance throughout the stretching process,Example 12: Waterproof treatment of textile

[0236] To develop and validate the waterproofing treatment used in the fabrication of the E-textile sensors disclosed herein, the following experiments were conducted

[0237] The application of textile-based electronics in real-life scenarios faces a significant challenge due to the inherent absorbency of textiles towards liquids. This characteristic makes the textile-based devices prone to damage by perspiration or water. To address this issue, we used 1 H,1 H,2H,2H- perfluorooctyltriethoxysilane (PFOTS), a fluorinated alkyl silane with a low surface free energy, which is suitable for forming anti-adhesive coatings on various materials, including glass, polymers, wafers, textiles, and evencellulose paper. An efficient one-step vapor- deposition technique has been developed to create a water-repellent protection layer on raw fabrics. To obtain high-quality self-assembled monolayer (SAM) of PFOTS, precise control of the vaporization temperature is crucial. The temperature was set at 90 °C, which is slightly below its flash point of 97 °C. It is worth noting that a droplet of 10 pL of PFOTS is sufficient for functionalizing a 4-inch textile patch. As PFOTS is heated on a hotplate, it gradually vaporizes and diffuses onto the target surface inside the reaction chamber. Furthermore, the hydroxyl groups present on textiles can react chemically with chlorosilanes, leading to the formation of surface- grafted siloxanes. This additional chemical reaction enhances the durability of the superhydrophobic coating. The entire silane functionalization process can be completed within minutes, making it a promising and cost- effective approach for large-scale device production and mass manufacturing.

[0238] To test the effectiveness of the waterproofing treatment, the resistance of a PFOTS-treated on-textile electrode sample was measured before and after being dipped in water. The measured resistance was essentially unchanged between pre-wetting (3326 ohm) and after wetting (3172 ohm).Example 13 Equivalent circuit model of the electrode-skin interface

[0239] To characterize the electric contact between the E-textile sensors disclosed herein and skin, the following experiments were conducted. Equivalent circuit models were developed and used to characterize the electrical interactions between the E-textile sensor and skin.

[0240] The electrode-skin interfaces and corresponding equivalent circuit models for various types of electrodes are illustrated below. For the conventional gelled electrodes (FIG. 25A), a half-cell potential (Ehc) is firstly formed at the electrochemical electrode-electrolyte interface. The bilayer structure of the electrode and gel could be modeled using a resistor (Rd) and a capacitor (Cd) in parallel. The gel, being rich in ions, could be denoted as a resistor (Rg). The skin is considered as a multilayer structure, consisting of epidermis, dermis, and subcutaneous tissue. The outmost layer of the epidermis is stratum corneum, which acts as a semipermeable layer composedof dead cells, resulting in a difference in ion concentrations and the presentence of potential (Ese). The epidermis layer could be modeled as a resistor (Re) in parallel with a capacitor (Ce). Lastly, the dermis and subcutaneous tissues consist of nerves, sweat glands, vessels, etc. can all be lumped together and modeled as a single resistor (Ru).

[0241] The elimination of gel usage could make the textile electronics wearing experience more comfortable and convenient, albeit at the expense of increased skin-electrode impedance and consequently deterioration in signal quality. The equivalent circuit models of on-textile electrodes and microfiber-on- textile electrodes operating in gel-free conditions are investigated. On one hand, in the absence of ionic gel, the textile electrodes are expected to exhibit substantial capacitive effects at the interface. On the other hand, the potential presence of moisture and sweat may contribute to the resistive effects. Therefore, the electrode-skin interface in gel-free conditions could be considered as a parallel combination of a resistor (Rd) and a capacitor (Cd). It can be observed that the on-textile electrodes have difficulty establishing conformal and reliable contact with the skin due to the inherent stiffness of textiles and the presence of hair, resulting in higher impedance and more pronounced motion artifacts (FIG. 25B). However, the utilization of microfiber- on-textile electrodes leads to a significant reduction in electrode-skin impedance, as shown in FIG. 25C, which can be attributed to the following factors:

[0242] 1 ) Numerous conductor microfibers evenly distributed on the skin surface substantially increase the contact area, thus reducing the electrodeskin impedance.

[0243] 2) The applied pressure has a substantial impact on the electrodeskin impedance. The presence of 3D microfibers enables easy penetration of hair, fills air gaps, and creates localized high pressure by slightly elevating and stretching the garment, without the need for intentional tightening or pressing of the electrode area.

[0244] 3) The electrode-skin impedance is inversely related to the skinhumidity and sweat, as they affect the conductivity and dielectric constant of the skin. The fluffy and absorbent microfibers can effectively absorb the perspiration and ambient moisture, securely adhering on the skin and reducing impedance.Example 14: Comparison of motion artifacts among different electrodes when used in dry condition

[0245] To evaluate the effects of motion artifacts on electrophysiological measurements obtained using the E-textile sensors as disclosed herein, the following experiments were conducted.

[0246] When using the commercial Ag / AgCI electrodes for biopotential recording, the electrodes are typically required to be firmly fixed at the desired location using 3M tapes and are thus usually capable of producing clean and stable signals when the subject is stationary. However, when subjected to body movements, the intense friction at the electrode-skin interface could result in significant motion artifacts that might interfere with the recorded signals. For the planar on-textile electrodes, poor contact with dry skin and high electrodeskin impedance led to noticeable motion artifacts caused by the subject's routine breathing activity even when the subject is stationary. In contrast, the 3D microfiber-on-textile electrodes exhibited good tolerance to motion artifacts and were able to record stable signals during both static and active movements. Moreover, despite the presence of small motion artifacts when the subject was in motion, distinguishable ECG characteristic peaks were clearly captured.

Claims

CLAIMSWhat is claimed is:1 . A wearable microfiber-based electrode for obtaining electrophysiological signals from a subject, the electrode comprising: a. an on-textile base electrode attached to an underlying textile material, the on-textile base electrode comprising a conductive polymer material; and b. a 3D conductive microfiber layer comprising a plurality of microfibers coated with the conductive polymer material, wherein the 3D conductive microfiber layer is attached the on-textile base electrode opposite the underlying textile material by an ionic binder layer. positioned between the on-textile base electrode and the 3D conductive microfiber layer.

2. The electrode of claim 1 , wherein the conductive polymer material comprises a conductive poly(3,4-ethylenedioxythiphene)-poly (styrene sulfonate) PEDOT:PSS polymer.

3. The electrode of claim 2, wherein the conductive polymer material further comprises a polymer additive, the polymer additive comprising polyethylene oxide) (PEG).

4. The electrode of claim 3, wherein the conductive polymer comprises the PDOT : PSS and the PEG in a weight ratio of up to 10: 1 .

5. The electrode of claim 4, wherein the conductive polymer comprises the PDOT : PSS and the PEO in a weight ratio of 5: 1 .

6. The electrode of claim 5, further comprising a hydrophobic coating comprising a self-assembled monolayer comprising perfluorooctyltrichlorosilane (PFOTS).

7. A wearable, wireless E-textile bioelectric recording system, the system comprising: a. at least two wearable microfiber-based electrodes printed on a textile substrate, each electrode comprising an on-textile base electrode attached to an underlying textile material, the on-textile base electrode comprising a conductive polymer material; and a 3D conductivemicrofiber layer comprising a plurality of microfibers coated with the conductive polymer material, wherein the 3D conductive microfiber layer is attached the on-textile base electrode opposite the underlying textile material by an ionic binder layer. positioned between the on- textile base electrode and the 3D conductive microfiber layer; b. a superhydrophobic fluorinated self-assembled monolayer deposited on the electrode surface; and c. a wireless data recording circuit operatively coupled to the at least two wearable microfiber-based electrodes to detect bioelectric signals from the electrodes.

8. The system of claim 7, wherein the conductive polymer material comprises a conductive poly(3,4-ethylenedioxythiphene)-poly (styrene sulfonate) PEDOT:PSS polymer.

9. The system of claim 8, wherein the conductive polymer material further comprises a polymer additive, the polymer additive comprising poly(ethylene oxide) (PEG).

10. The system of claim 9, wherein the conductive polymer comprises the PDOT : PSS and the PEG in a weight ratio of up to 10: 1 .11 .The system of claim 10, wherein the conductive polymer comprises the PDOT: PSS and the PEO in a weight ratio of 5:1 .

12. The system of claim 11 , further comprising a hydrophobic coating comprising a self-assembled monolayer comprising perfluorooctyltrichlorosilane (PFOTS).

13. A method to monitor the health and physiology of a subject, the method comprising having the subject wear a smart garment containing the E-textile bioelectric recording system of any previous claim on or in a region of interest and recording bioelectric signals of the subject with the E-textile system.

14. The method of claim 13, wherein the system performs real-time multimodal electrophysiological signal monitoring.

15. The method of claim 14, wherein the real-time signal monitoring is selected from an electrocardiogram (ECG), an electromyography (EMG), and any combination thereof.

16. The method of claim 15, wherein the real-time monitoring is carried out during strenuous exercise by a subject.

17. The method of claim 16, wherein the ECG is a maternal ECG, and the EMG is a uterine EMG of a pregnant mother.

18. The method of claim 17, wherein the maternal ECG and the uterine EMG contributes to uterus-relevant physiological assessments and early detection of abnormal uterine contraction patterns and birth-related risks.

19. A method to fabricate an E-textile bioelectric recording device, the method comprising: a. printing at least two 3D microfiber-based electrodes onto a textile, b. vaporizing a self-assembled monolayer of a superhydrophobic monolayer onto the electrodes, and a. connecting a custom-designed motion-artifact cancelling wireless data recording circuit to the electrodes to record bioelectric signals.

20. The method of claim 19, wherein the electrodes are composed of a conducting polymer.

21. The method of claim 20, wherein the conducting polymer is PEDOT:PSS.

22. The method of claim 21 , wherein the superhydrophobic monolayer is composed of perfluorooctyltrichlorosilane (PFOTS).

23. The method of claim 22, wherein the device is equipped with at least two recording channels by connecting at least two electrodes to the recording circuit.

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