Textile electrode, method of fabrication and application(s) thereof

Abstract

The present disclosure provides a textile electrode for biomonitoring, enabling consistent and efficient monitoring of health while operating in contact with a subject's body. Particularly provided in the present disclosure is a textile electrode comprising a conductive textile coated with a thickness-controlled layer of a non-conductive elastomeric polymer. The said textile electrode is characterized by advantages such as retained conductivity of the conductive textile, structural stability strong adherence to skin, reusability after multiple cycles of washing and ease of fabrication.

Description

[0001]“TEXTILE ELECTRODE, METHOD OF FABRICATION AND APPLICATION(S) THEREOF” TECHNICAL FIELD The present disclosure relates to the fields of electronic sensor technologies, and their incorporation into textiles. The present disclosure particularly provides a reusable textile electrode for biomonitoring, enabling consistent and efficient monitoring of health while operating in contact with a subject’s body. BACKGROUND OF THE DISCLOSURE Electrocardiography (ECG) measures the electrical activity of the heart. The electric currents from cardiac muscle contractions produce detectable voltages that reflect depolarization and repolarization of cells [1]. It captures the potential difference across the skin's surface using conductive electrodes. ECG electrodes can be categorized as wet or dry. Wet electrodes have an electrolytic gel to enable signal transmission to the electrode. One reusable type is clamp / bulb electrodes, but they do not adhere well on skin and are typically used for short term monitoring. Other types of electrodes usually consist of a foam sticker, with an Ag / AgCl metal contact and a solid gel at the center. The solid gel electrodes provide good skin adhesion and signal quality, but they are disposable and not reusable. The presence of the gel in wet electrodes show that resistive coupling dominates at the electrode-electrolyte junction [2]. The conductive gel also hydrates the skin, helping reduce skin-electrode impedance. While wet electrodes can provide high quality signals for brief recordings, they are not optimal for long- term monitoring because the conductive gel dries out over time, resulting in declining signal quality. Dry electrodes address wet electrode challenges like long term usability, reusability, and comfort but their lack of gel leads to impedance issues, drift, and poor repeatability. In dry electrodes, the lack of electrolytic gel means that there are voids / air pocket at the junction between the electrode and skin that results in capacitive coupling or high skin impedance. This results in signals with poor signal to noise ratio (SNR) [3]. Textile-based dry electrodes are promising due to the ability to conform to the skin, breathability, and flexibility. These electrodes can be developed by using smart textiles obtained by integrating conductive fibers or threads onto fabric through weaving or knitting. Other existing methods for textile electrodes involve complex and multi-step fabrication methods like electronic printing where conductive inks are deposited on textile substrates, embroidery, sewing conductive yarn onto textile substrates or growing nanostructures such as graphene or CNT on textile. These methods have high production costs and involve complexity, making them unsuitable for mass production. Particularly, embroidery involves sewing conductive yarns onto a textile substrate, which requires careful selection of materials that can withstand the embroidery process without tearing, stretching, or deforming. This limitation restricts the range of suitable materials for embroidered electrodes. Furthermore, the durability of embroidered electrodes may be compromised by repeated washing, as the conductive yarns can loosen and detach from the substrate. Screen printing, while widely used for mass production in the textile industry, requires careful selection and formulation of conductive inks to prevent mechanical cracking, which can affect the signal being measured and compromise the electrode's performance and durability. There is therefore a need in the art for efficient dry electrodes that address the aforementioned drawbacks associated with such electrodes. SUMMARY OF THE DISCLOSURE Addressing the aforesaid requirement in the art, provided herein is a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns. In some embodiments, the conductive textile comprises a knitted or woven fabric coated with a conductive substance. In some embodiments, the knitted or woven fabric is a porous stretchable fabric. In some embodiments, the knitted or woven fabric comprises polyamide, elastane, cotton and viscose or any combination thereof; and / or wherein the conductive substance is selected from a group comprising metal(s) and conductive carbon-based material(s) or a combination thereof. In some embodiments, non-limiting examples of the metal(s) include silver, gold, copper and stainless steel or any combination thereof. The said coating with the conductive substance(s) renders the textile conductive. In some embodiments, the conductive textile has thickness ranging from about 0.40 mm to about 0.70 mm. In some embodiments, the insulating elastomeric polymer comprises polymer(s) selected from a group comprising silicone(s) and polydimethylsiloxane (PDMS) or a combination thereof. Without intending to be limited by theory, the insulating elastomeric polymer stabilizes shape of the conductive textile and enables stable skin adhesion. In some embodiments, the insulating elastomeric polymer confers reusability / durability to the textile electrode such that the electrical and mechanical properties of the textile electrode have minimal deterioration after multiple cycles of washing. Thus, in some embodiments, the textile electrode is a reusable textile electrode that retains its shape and conductivity after one or more cycles of washing and / or exposure to cleaning agents. In some embodiments, reduction in SNR after multiple cycles of washing is limited to about 10% to about 25%. In some embodiments, reduction in SNR after at least 10 cycles of washing is limited to about 10% to about 25%. Further provided herein is a method of fabricating the textile electrode as described above, comprising: casting the insulating elastomeric polymer on the conductive textile to obtain an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns over the conductive textile. In some embodiments, the casting is performed at a temperature of about 20°C to about 25°C. In some embodiments, the casting step is followed by curing the insulating elastomeric polymer cast on the conductive textile. In some embodiments, the curing is performed at a temperature of about 50°C to about 80°C; and / or wherein the curing is performed for about 8 minutes to about 25 minutes. The present disclosure also provides a device for real-time biofeedback comprising the textile electrode as described above. In some embodiments, the device is an electrocardiography (ECG) device for Heart Rate Variability Monitoring (HRV). In some embodiments, the device further comprises an additional textile electrode for respiration monitoring to measure Respiratory inductance plethysmography (RIP). In a non-limiting embodiment, the device is in the form of a skin patch. Further envisaged herein is an ECG monitoring kit comprising the textile electrode or the device as claimed as described above. The said kit, in some embodiments, may further comprise one or more of an instruction manual, a signal acquisition module, signal processing unit, an ECG evaluation board, wire(s), an accelerometer and soft skin adhesives for additional support. The present disclosure also provides use of the textile electrode, or the device as described above in monitoring electrical activity of the heart. BRIEF DESCRIPTION OF THE ACCOMPANYING FIGURES In order that the disclosure may be readily understood and put into practical effect, reference will now be made to exemplary embodiments as illustrated with reference to the accompanying figures. The figures together with detailed description below, are incorporated in and form part of the specification, and serve to further illustrate the embodiments and explain various principles and advantages, where: Figure 1 depicts (a) different types of wet ECG electrodes in the art; (b) design of the textile electrode of the present disclosure. Figure 2 depicts (a) textile electrodes without any backing layer after one wash; (b) Textile electrodes with a 500-micron backing layer after multiple washes. Figure 3 depicts textile surface after silicone coating indicating the voids on the backside of the textile. Figure 4 depicts positioning of the textile electrodes. Figure 5 depicts comparison of raw ECG signals between the conventional hydrogel electrode and the textile electrode of the present disclosure (a) Unwashed (b) After washing for 6 cycles and filtered signal of (c) Unwashed (d) After washing for 6 cycles. Figure 6 depicts an analysis comparing the signal-to-noise ratio (SNR) across three different sets of electrodes after multiple cycles of handwashing. Figure 7 depicts an analysis comparing the SNR across three different sets of electrodes after cleaning with IPA. Figure 8 depicts the correlation between thickness of the polymer layer and the textile electrode’s conductivity. Figure 9 depicts the textile electrode having a polymer layer thickness of about 200μm before and after use. Figure 10 depicts comparison between ECG Signal of Textile-Silicone (Ecoflex) vs Textile- polydimethylsiloxane (PDMS) (Unfiltered Data). Figure 11 depicts comparison between ECG Signal of Textile-Silicone (Ecoflex) vs Textile- polydimethylsiloxane (PDMS) (Filtered Data). DETAILED DESCRIPTION OF THE INVENTION Addressing the aforesaid need pertaining to the requirement of dry electrodes, the present disclosure provides a reusable textile-based electrode that is easy to fabricate. Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the technical domain or art to which this invention belongs. Although any methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present invention, the preferred methods and materials are now described. A number of terms are used throughout the specification with the following definitions provided for convenience. General definitions The term “dry electrode” as used throughout the present disclosure refers to non-invasive electrodes which are applied directly to the skin and which do not require any gel or saline liquid for ensuring conductivity. The term “textile electrode” as used throughout the present disclosure refers to the electrode of the present disclosure composed of a conductive textile and an elastomeric polymer coating, wherein the elastomeric polymer coating is non-conductive in nature and provides a supportive backing to the textile electrode. The term “conductive textile” refers to the fabric comprised in the textile electrode, which is characterized by coating or treatment with a conductive substance to render it conductive. The phrase “coated with a conductive substance” refers to the introduction of a metallic or other conductive component such as but not limited to conductive carbon-based materials in the fabric in the form of a plating or coating of the fabric or individual yarns in the fabric that is knitted or woven, to yield a conductive textile. As used herein, the term “biomonitoring” refers to the long-term or short-term acquisition and recording of physiological parameters such as but not limited to Electrocardiography (ECG) signal, heart rate variability (HRV), Respiratory inductance plethysmography (RIP) and breathing of a monitored subject. The term “Biofeedback” refers to controlling involuntarily processes in the body such as breathing, Heart rate blood pressure etc. from the feedback of sensors. Reference to “insulating elastomeric polymer”, “elastomeric coating” and “non-conductive elastomeric coating” or “polymer coating” is with respect to the thickness controlled non- conductive polymeric coating on the conductive textile in the textile electrode of the present disclosure. As used herein, the term “subject” is a vertebrate, such as a mammal, such as a human. Mammals include, but are not limited to, humans, livestock, athletic animals, pets and the like. As used herein, the term “comprising” when placed before the recitation of steps in a method means that the method encompasses one or more steps that are additional to those expressly recited, and that the additional one or more steps may be performed before, between, and / or after the recited steps. For example, a method comprising steps a, b, and c encompasses a method of steps a, b, x, and c, a method of steps a, b, c, and x, as well as a method of steps x, a, b, and c. Furthermore, the term “comprising” when placed before the recitation of steps in a method does not (although it may) require sequential performance of the listed steps, unless the content clearly dictates otherwise. For example, a method comprising steps a, b, and c encompasses, for example, a method of performing steps in the order of steps a, c, and b, the order of steps c, b, and a, and the order of steps c, a, and b, etc. With respect to the use of substantially any plural and / or singular terms herein, those having skill in the art can translate from the plural to the singular and / or from the singular to the plural as is appropriate to the context and / or application. The various singular / plural permutations may be expressly set forth herein for sake of clarity. The suffix “(s)” at the end of any term in the present disclosure envisages in scope both the singular and plural forms of said term. As used in this specification and the appended claims, the singular forms “a,” “an” and “the” includes both singular and plural references unless the content clearly dictates otherwise. For example, the term “inserted at a position” as used herein in reference to a polypeptide sequence refers to insertion at one or more (such as one, two, three, etc.) amino acid positions in the polypeptide sequence. The use of the expression ‘at least’ or ‘at least one’ suggests the use of one or more elements or ingredients or quantities, as the use may be in the embodiment of the disclosure to achieve one or more of the desired objects or results. As such, the terms “a” (or “an”), “one or more”, and “at least one” can be used interchangeably herein. Numerical ranges stated in the form ‘from x to y’ include the values mentioned and those values that lie within the range of the respective measurement accuracy as known to the skilled person. If several preferred numerical ranges are stated in this form, of course, all the ranges formed by a combination of the different end points are also included. The terms “about” or “approximately” as used herein when referring to a measurable value such as a parameter, an amount, a temporal duration, and the like, are meant to encompass variations of and from the specified value, such as variations of + / -10% or less, + / -5% or less, + / -1% or less, and + / -0.1% or less of and from the specified value, insofar such variations are appropriate to perform in the disclosed invention. It is to be understood that the value to which the modifier “about” or “approximately” refers is itself also specifically, and preferably, disclosed. As used herein, the terms “include” (any form of “include”, such as “include”), “have” (and “have”), “comprise” etc. any form of “having”, “including” (and any form of “including” such as “including”), “containing”, “comprising” or “comprises” are inclusive and will be understood to imply the inclusion of a stated element, integer or step, or group of elements, integers or steps, but not the exclusion of any other element, integer or step, or group of elements, integers or steps As regards the embodiments characterized in this specification, it is intended that each embodiment be read independently as well as in combination with another embodiment. For example, in case of an embodiment 1 reciting 3 alternatives A, B and C, an embodiment 2 reciting 3 alternatives D, E and F and an embodiment 3 reciting 3 alternatives G, H and I, it is to be understood that the specification unambiguously discloses embodiments corresponding to combinations A, D, G; A, D, H; A, D, I; A, E, G; A, E, H; A, E, I; A, F, G; A, F, H; A, F, I; B, D, G; B, D, H; B, D, I; B, E, G; B, E, H; B, E, I; B, F, G; B, F, H; B, F, I; C, D, G; C, D, H; C, D, I; C, E, G; C, E, H; C, E, I; C, F, G; C, F, H; C, F, I, unless specifically mentioned otherwise. Disclosure Conventional electrodes, also called ‘wet electrodes’ (Figure 1), use liquid or solid gels that can be uncomfortable, dehydrate over time causing signal loss, are not reusable and are also not suitable for long term biomonitoring. Dry electrodes address challenges associated with wet electrodes like long term usability, reusability, and comfort but their lack of gel often leads to impedance issues, drift, and poor repeatability. It is thus an object of the present disclosure to provide dry electrodes that obviate the reliance on such gels to enable, simplify and improve the efficiency of long-term biomonitoring. Non-limiting examples of such biomonitoring applications include electrocardiography (ECG), Heart Rate Variability Monitoring (HRV) and the like. Textiles that are knitted or woven from conductive yarns alone are often unreliable as electrodes because they lose their shape without a supporting layer. This is depicted in Figure 2(a). It is therefore another object of the present disclosure to provide a textile electrode that addresses the drawbacks of previously known textile-based electrodes and is characterized by shape retention, signal fidelity and skin interface stability. Accordingly, addressing drawbacks in the art relating to dry electrodes and specifically textile- based electrodes, the present disclosure provides a reusable textile electrode comprising a layer of insulating elastomer coated over a conductive textile. Critically, albeit being of insulating nature, the elastomeric polymer coating present on the textile electrode is engineered to a thickness that avoids insulating the conductive textile layer underneath completely. The controlled thickness of the coating opens small voids in the conductive textile that retain conductivity for capturing ECG signals. This delicate balance provides a backing layer in the form of the insulating elastomeric polymer coating to stabilize the shape of the underlying conductive textile while preserving its conductivity. Textile electrode Particularly, the present disclosure provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns. In some embodiments, the conductive textile comprises a knitted or woven fabric coated with a conductive substance. In some embodiments, the knitted or woven fabric is a porous stretchable fabric. Accordingly, in some embodiments, the conductive textile is a porous stretchable fabric coated with a conductive substance. The porous nature of the conductive textile allows it to partially absorb the thin insulating polymer coating. Further, the voids formed in the coating due to the rough surface of the underlying textile preserves the inherent conductivity of the textile, enabling high signal to noise ratio (SNR) signals. Without intending to be limited by theory, the porosity of the textile and application of the insulating elastomeric polymer coating allows both tight skin adhesion as well as maintained surface voids in the conductive textile for continued conductivity. Said properties of the textile electrode are particularly relevant and important for biosensing applications such as electrocardiography (ECG or EKG). In some embodiments, the polymer layer thickness may be adjusted within the above defined range of about 300 microns to about 600 microns based on the non-conductive textile's porosity to ensure uniformity and maintain conductivity. Additionally, the stretchable nature of the conductive textile allows it to conform to the skin surface, thereby leading to better adherence which is crucial for the said intended applications. In some embodiments, the conductive textile comprises polyamide, elastane, cotton and viscose or any combination thereof. In a preferred embodiment, the conductive textile comprises a combination of polyamide and elastane. In a non-limiting embodiment, the conductive textile comprises about 72% to about 95% of polyamide and about 5% to about 28% of elastane. In some embodiments, the conductive textile comprises about 72%, about 73%, about 74%, about 75%, about 76%, about 77%, about 78%, about 79%, about 80%, about 81%, about 82%, about 83%, about 84%, about 85%, about 86%, about 87%, about 88%, about 89%, about 90%, about 91%, about 92%, about 93%, about 94% or about 95% of polyamide and about 5%, about 6%, about 7%, about 8%, about 9%, about 10%, about 11%, about 12%, about 13%, about 14%, about 15%, about 16%, about 17%, about 18%, about 19%, about 20%, about 21%, about 22%, about 23%, about 24%, about 24%, about 25%, about 26%, about 27% or about 28% of elastane. In some embodiments, the conductive textile comprises about 79% polyamide and about 21% elastane. In some embodiments, the conductive textile comprises about 94% polyamide and about 6% elastane. In some embodiments, the conductive textile comprises about 79% polyamide and about 21% elastane. While the said conductive textiles may differ in their stretchability due to the varying ratios of polyamide to elastane, they share similar properties, such as porosity and conductivity when coated with a conductive substance. In preferred embodiment, the conductive textile comprises about 78% polyamide and about 22% elastane. In some embodiments, the textile comprises the above identified fabric(s) coated with a conductive substance selected from a group comprising metal(s) and conductive carbon-based material(s) or a combination thereof. In some embodiments, non-limiting examples of the metal(s) include silver, gold, copper and stainless steel or any combination thereof. The said coating with the conductive substance(s) renders the textile conductive. In a preferred embodiment, the textile is coated with silver as the conductive substance. In some embodiments, the coating with the conductive substance is achieved by employing a knitted fabric partially or completely coated with a conductive substance. In some embodiments, the individual yarns forming the knitted or woven fabric may be coated with the conductive substance. In a preferred embodiment, the individual yarns forming the knitted or woven fabric may be silver plated. In some embodiments, the conductive textile comprises polyamide, elastane, cotton and viscose or any combination thereof coated with a conductive substance selected from a group comprising metal(s) and conductive carbon-based material(s) or a combination thereof. In a preferred embodiment, the conductive textile comprises polyamide and elastane coated with silver. In other words, in some embodiments, the conductive textile is a silver-plated polyamide-elastane fabric. Accordingly, in some embodiments, the present disclosure provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein the conductive textile is a porous stretchable fabric. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein the conductive textile is a polyamide-elastane fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein the conductive textile is a silver-plated polyamide-elastane fabric. In some embodiments, thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm. In some embodiments, thickness of the conductive textile ranges from about 0.40 mm to about 0.50 mm, 0.50 mm to about 0.60 mm, 0.50 mm to about 0.70 mm, about 0.40 mm, about 0.45 mm, about 0.50 mm, about 0.55 mm, about 0.60 mm, about 0.65 mm or about 0.70 mm. In a preferred non-limiting embodiment, thickness of the conductive textile is about 0.55 mm. Accordingly, in some embodiments, the present disclosure provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a porous stretchable fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a polyamide-elastane fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a silver-plated polyamide-elastane fabric. Some textile sensors in the art comprise a fabric paired with conductive polymers to impart stability and conductivity. Conductive polymers, however, have to be prepared from scratch, thus requiring complex fabrication processes to integrate the polymer and conductive components in just the right proportions, potentially leading to issues in terms of cost of production. The present disclosure instead provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer. Use of the insulating elastomeric polymer in the present disclosure ensures a single step fabrication and makes it easy for large scale manufacturing as opposed to conductive polymers which require careful and complex fabrication, ensuring consistency across batches, before application to textile. In some embodiments, the insulating elastomeric polymer is one that shows the property of skin adhesion. The said property of skin adhesion is of specific significance since one of the intended applications of the textile electrode is as a biomonitoring sensor such as an ECG sensor, wherein the polymer coated side of the textile electrode faces or adheres to the skin. In some embodiments, the insulating elastomeric polymer may be selected such that it has similar Youngs Modulus as skin, which helps in it have skin-like properties and adherence to skin upon application. In some embodiments, the insulating elastomeric polymer comprises polymer(s) selected from a group comprising silicone(s) and polydimethylsiloxane (PDMS) or a combination thereof. Use of other polymers that confer similar property of skin adhesion to the electrode of the present disclosure is envisaged in the scope of the present disclosure. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns; wherein the conductive textile comprises a combination of polyamide and elastane; and / or wherein the insulating elastomeric polymer comprises silicone(s). In some embodiments, the conductive textile comprises the insulating elastomeric polymer layer on one side of the conductive textile, wherein the said side is used for skin contact or skin adhesion. As mentioned above, the said insulating elastomeric polymer layer is employed at a controlled thickness such that the coating does not lead to loss of conductivity of the conductive textile. Specifically, the insulating elastomeric polymer layer over the conductive textile is of a thickness ranging from about 300 microns to about 600 microns, as mentioned above. In some embodiments, the insulating elastomeric polymer layer is of a thickness of about 300 microns to about 400 microns, about 350 microns to about 500 microns, about 400 microns to about 500 micros, about 300 microns, about 320 microns, about 340 microns, about 360 microns, about 380 microns, about 400 microns, about 420 microns, about 440 microns, about 460 microns, about 480 microns, about 500 microns, about 520 microns, about 540 microns, about 560 microns, about 580 microns or about 600 microns. In a preferred embodiment, the insulating elastomeric polymer layer on the conductive textile is of a thickness of about 400 microns to about 600 microns. In some preferred embodiments, the insulating elastomeric polymer layer on the conductive textile is of a thickness of about 450 microns to about 550 microns. In a non-limiting, preferred embodiment, the insulating elastomeric polymer layer on the conductive textile is of a thickness of about 500 microns. Accordingly, in some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns. In some embodiments, the present disclosure provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer of thickness ranging from about 450 microns to about 550 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a porous stretchable fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; wherein the conductive textile is a porous stretchable fabric coated with a conductive substance; and wherein the insulating elastomeric polymer layer is composed of silicone(s). In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a polyamide-elastane fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns; thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a silver-plated polyamide-elastane fabric. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 450 microns to about 550 microns; thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; wherein the conductive textile is a silver-plated polyamide-elastane fabric; and wherein the insulating elastomeric polymer layer is composed of silicone(s). In some embodiments, the present disclosure provides a textile electrode comprising a conductive textile coated with an insulating elastomeric polymer of thickness of about 500 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a porous stretchable fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; wherein the conductive textile is a porous stretchable fabric coated with a conductive substance; and wherein the insulating elastomeric polymer layer is composed of silicone(s). In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; and wherein the conductive textile is a polyamide-elastane fabric coated with a conductive substance. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein thickness of the conductive textile ranges from about 0.40 mm to about 0.70 mm; wherein the conductive textile is a polyamide-elastane fabric coated with a conductive substance; and wherein the insulating elastomeric polymer layer is composed of silicone(s). In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein the conductive textile is a silver-plated polyamide-elastane fabric. In some embodiments, the textile electrode comprises a conductive textile coated with an insulating elastomeric polymer layer of thickness of about 500 microns; wherein the conductive textile is a silver-plated polyamide-elastane fabric; and wherein the insulating elastomeric polymer layer is composed of silicone(s). Controlled thickness of the insulating polymer coating on the conductive textile, as mentioned above, balances conductivity and stability. The insulating elastomeric polymer stabilizes the shape of the conductive textile and enables stable skin adhesion by preventing textile wrinkling and deformation that can reduce electrode contact. This allows robust contact with the skin of the subject for quality signal detection. The textile electrode of the present disclosure thus provides signal to noise ratio (SNR) higher than conventional ECG electrodes that for instance rely on foam sticker for adhesion. The insulating elastomeric polymer coating further ensures that the textile’s original shape over time is retained avoiding distortion even after repeated use and washing. This shape retention thus enables consistent performance of the textile electrode as a reusable electrode. The combined advantages of shape retention, signal fidelity, and skin interface stability make the textile electrode of the present disclosure uniquely suited for reliable long-term biomonitoring in wearable applications. The elastomeric layer, thus, primarily provides three key benefits: a) the textile electrode maintains its shape, electrical and mechanical properties after repeated washing cycles, and b) the elastomeric polymer with inherent stickiness helps achieve tight skin contact for high SNR signals and c) the electrical conductivity of the textile electrodes is intact for detecting the ECG measurements. Overall, the insulating elastomeric polymer employed in the textile of the present disclosure confers to it: - Mechanical Properties – Support to the conductive textile in the form of a stable backing layer; - Adhesive Properties - The polymer has an inherent stickiness that helps in adhering to the skin well; and - Retained Conductivity - Creation of the voids due to absorption of the polymer before curing ensures openings to maintain the conductivity of the underlying textile. In a non-limiting embodiment, the textile electrode of the present disclosure may comprise the insulating elastomeric polymer one or more sides of the conductive textile, wherein while the side for skin contact has controlled thickness of the insulating elastomeric polymer layer, the other side(s) may or may not have controlled thickness of the insulating elastomeric polymer layer. In some embodiments, the textile electrode of the present disclosure is a reusable textile electrode which retains the textile’s original shape, avoiding distortion after repeated use and washing. The ability to retain high signal quality, stable adhesion, and mechanical robustness after repeated washing and abrasion cycles, unlike gel electrodes that lose function after single use, allows for long-term ECG monitoring applications. Figure 2b shows the developed textile- polymer electrode after 10 washes where there is no deformation or significant change in the electrical or mechanical properties of the electrodes. In some embodiments, the reusable textile electrode also retains its shape and conductivity after exposure to cleaning agents such as but not limited to alcohol. Accordingly, in some embodiments, the provided herein is a reusable textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns, wherein the reusable textile electrode retains its shape and conductivity after one or more cycles of washing and / or exposure to cleaning agents. In some embodiments, reduction in SNR after multiple cycles of washing is limited to about 10% to about 25%. In some embodiments, reduction in SNR after at least 10 cycles of washing is limited to about 10% to about 25%. Overall, the textile electrode of the present disclosure enables durable and high-fidelity ECG monitoring through the synergistic integration of a robust conductive textile and a skin- conforming adhesive polymer layer that helps to lower the skin impedance of the electrodes. Without intending to be limited by theory, due to the skin conforming effect of the textile electrode of the present disclosure, there is an increase in the effective contact area which results in a reduction in the electrode-skin impedance. The textile electrode employs a flexible combination of textiles and polymers that conforms to the skin for comfort and breathability, thereby providing an unobtrusive wearable solution for long-term biomonitoring. Method of fabricating the textile electrode The hybrid textile-elastomer approach as described above yields a reusable, high-performing textile electrode without complex production steps. Without intending to be limited by theory, the elastomeric coating on the textile electrode is achieved using a single step casting method. Accordingly, provided herein is a method of fabricating the textile electrode as described above, comprising: casting the insulating elastomeric polymer on the conductive textile to obtain an insulating elastomeric polymer layer of thickness ranging from about 300 to about 600 microns over the conductive textile. In some embodiments, the casting is by manual, automated or mechanized application. Without being restricted by the language used above, the casting step as mentioned above may comprise application of the insulating elastomeric polymer to the conductive textile or application of the conductive textile to the insulating elastomeric polymer. In a non-limiting embodiment, the casting comprises coating a layer of the insulating elastomeric polymer on a glass plate with an applicator and then placing the textile on top. In some embodiments, in order to facilitate the casting, the polymer may be poured on a glass plate / suitable rigid plate. Then, a manual applicator such as a doctor blade with screws to adjust the height may be used to slide / move the blade which results in a layer of the elastomeric polymer. The height of the manual applicator can be adjusted which results in a polymer of the required thickness. In some embodiments, the casting is performed at room temperature. In some embodiments, the casting is performed at a temperature of about 20°C to about 25°C. In some embodiments, the casting step is followed by curing the insulating elastomeric polymer cast on the conductive textile. In some embodiments, the insulating elastomeric polymer may thus comprise or may be combined with a curing agent during its preparation to enable efficient curing. In some embodiments, the curing is performed at a temperature of about 50°C to about 80°C. In some embodiments, the curing is performed at a temperature of about 50°C to about 60°C, about 60°C to about 70°C, about 70°C to about 80°C, about 50°C, about 55°C, about 60°C, about 65°C, about 70°C, about 75°C or about 80°C. In a preferred embodiment, the curing is performed at a temperature of about 70°C. In some embodiments, the curing is performed for about 8 minutes to about 25 minutes based on the size of the conductive textile. In some embodiments, the curing is performed for about 8 minutes to about 13 minutes, 13 minutes to about 23 minutes, 18 minutes to about 25 minutes, about 8 minutes, about 10 minutes, about 12 minutes, about 14 minutes or about 16 minutes, about 18 minutes, about 20 minutes, about 22 minutes, about 24 minutes or about 25 minutes. In a preferred embodiment, the curing is performed for about 15 minutes. In some embodiments, the curing is performed at a temperature of about 50°C to about 80°C; and / or the curing is performed for about 8 minutes to about 25 minutes. Thus, in some embodiments, the method of fabricating the textile electrode as described above, comprises: casting the insulating elastomeric polymer on the conductive textile at a thickness ranging from about 300 to about 600 microns; and curing the conductive textile casted with the insulating elastomeric polymer to obtain the textile electrode. In some embodiments, the method of fabricating the textile electrode as described above, comprises: casting the insulating elastomeric polymer on the conductive textile at of thickness ranging from about 300 to about 600 microns; and curing the conductive textile casted with the insulating elastomeric polymer at a temperature of about 70°C, for about 15 minutes to obtain the textile electrode. Without intending to be limited by theory, curing the textile cast with polymer may form a mat of the conductive textile coated with layer(s) of insulating elastomeric coating with not limitation on dimensions. The obtained insulating elastomeric polymer coated conductive textile may then be cut as per desired dimensions to obtain the textile electrode. As mentioned above, unlike disposable gel electrodes that lose function after single use, the textile electrode of the present disclosure maintains high signal fidelity over extended monitoring without degradation. The complementary combination of the conductive textile base and multifunctional insulating elastomeric polymer coating enables both strong adhesive properties and stable conductivity through an efficient single step fabrication technique. This results in a durable, affordable, and sensitive electrode fit for biomonitoring applications such as but not limited to ECG that can be mass produced cost-effectively. Device comprising the textile electrode The present disclosure further provides a device for real-time Heart Rate Variability Biofeedback comprising the textile electrode as described above. Accordingly, in some embodiments, provided herein is a device for real time biosensing or biomonitoring comprising the textile electrode of the present disclosure. In some embodiments, the device is an electrocardiography (ECG) device for Heart Rate Variability Monitoring (HRV). Unlike gel electrodes that lose function after single use, the device of the present disclosure provides an inventive solution for long-term ECG monitoring applications. In some embodiments, the device may further comprise additional electrode(s) for biomonitoring applications such as but not limited to electromyography, pressure sensing, strain sensing and respiration monitoring. In some embodiments, the device may further comprise additional electrode(s) to measure Respiratory Inductance plethysmography (RIP). RIP uses piezoresistive sensors in addition to the textile electrode of the present disclosure to detect changes in the local strain of the user's rib cage and abdomen during breathing cycles, enabling measurement of respiration rate and volume. The combination of HRV score and RIP helps analyze and understand the performance of the autonomous nervous system of the body which corelates to how the body responds to stress. The present disclosure particularly addresses the need for early, proactive chronic stress management by providing people the means to monitor and mitigate their stress effectively, ahead of the curve and before negative health impacts arise. In some embodiments, the device may comprise one or more of a signal acquisition module, signal processing unit and an ECG evaluation board. The data may be collected from such modules and transferred to a computer / phone. In a non-limiting embodiment, these modules may be present on a circuit board such as but not limited to a flex PCB beneath the patch to collect the ECG and / or respiration data. In a non-limiting embodiment, the textile electrode and / or the device may comprise wiring to enable connection of the textile electrode with the ECG evaluation board. In a non-limiting embodiment, the said wiring may be composed of specialized yarns coated with material(s) such as but not limited to Thermoplastic polyurethane (TPU) for shielding from noise. In some embodiments, the device may further comprise an accelerometer for movement measurements and signal artifact correction. In some embodiments, the device is in the form of a skin patch. The said skin patch, in a non- limiting embodiment, may comprise one or more further components such as but not limited to electronics as described above, a further backing layer, a polymeric film and soft skin adhesive(s). In some embodiments, the said skin patch is designed as a bio-sticker or sticker-like device that attaches directly to the skin. Kit Further envisaged in the present disclosure is an ECG monitoring kit comprising the textile electrode of the present disclosure. In some embodiments, the kit may comprise the textile electrode in the form of the textile electrode as such or in the form of a device or skin patch into which the textile electrode has been incorporated, as described above. In some embodiments, the kit may optionally comprise further components or instruments such as but not limited to a signal acquisition module, signal processing unit, an ECG evaluation board, wire(s), and an accelerometer. Accordingly, envisaged herein is an ECG monitoring kit comprising the textile electrode or device of the present disclosure, optionally in combination with one or more of an instruction manual, a signal acquisition module, signal processing unit, an ECG evaluation board, an accelerometer, wire(s) and / or soft skin adhesives for additional support. The kit, in some embodiments, may comprise further electronics or means for facilitating efficient signal acquisition, filtering and evaluation. In some embodiments, the kit may further comprise an instruction manual to provide instructions for using the textile electrode or device for biomonitoring applications and / or optionally, a soft skin adhesive for additional support in skin adherence. In addition to passive measurement, the device, skin patch and / or the kit as described above also provides real-time biofeedback to actively increase heart rate variability using resonant breathing techniques. To determine the resonant frequency unique to each user, the device, skin patch and / or the kit may incorporate a textile electrode for respiratory inductance plethysmography (RIP). Intended Application The textile electrode of the present disclosure or the device comprising the textile electrode finds application in monitoring electrical activity of the heart. In some embodiments, the textile electrode or the device comprising the textile electrode is intended for biomonitoring applications such as but not limited to Electrocardiography (ECG), heart rate variability (HRV), Respiratory inductance plethysmography (RIP) and breathing of a subject. Accordingly, the present disclosure provides use of the textile electrode or the device comprising the textile electrode as described above in ECG monitoring. Also envisaged herein is use of the textile electrode or the device comprising the textile electrode in monitoring one or more of Electrocardiography (ECG), heart rate variability (HRV), Respiratory inductance plethysmography (RIP) and breathing of a subject. Advantages of the textile electrodes of the present disclosure - Preservation of the inherent conductivity of the underlying conductive textile, enabling high SNR signals. - The elastomeric layer adheres to skin without the wrinkling or deformation that untreated textiles often undergo. This allows robust skin contact for signals with minimal attenuation. - Formation of voids on the textile electrode when the applied elastomeric polymer coating has an optimal thickness which stabilizes the textile shape while preserving textile conductivity. - Antimicrobial nature of the textile electrode due to the metallized conductive textile . - Reusability after multiple cycles of washing and cleaning. - Ease of fabrication by a single-step casting process. - The unobtrusive wearable solution makes stress management effortless and effective. Unlike rigid traditional sensors or optical heart rate variability (HRV) measurement devices, the textile electrode device uses a flexible combination of textiles and polymers that conforms to the skin for comfort and breathability during long-term monitoring. It utilizes electrocardiography (ECG), the gold standard for HRV measurement, for robust, noise-free signal acquisition even during movement and exercise. - While algorithms detect the user's heart rate variability dropping below their optimal range, tailored haptic and visual cues guide them through exercises to rebalance into healthy zones. Regular use of the electrode, device and / or kit trains resilience against stress and empowers people to take control of their health. The result is reduced anxiety, improved focus, and better quality of life. The subject matter of embodiments of the present invention is described here with specificity to meet statutory requirements, but this description is not necessarily intended to limit the scope of the disclosure. The disclosed subject matter may be embodied in other ways, may include different elements or steps, and may be used in conjunction with other existing or future technologies. This description should not be interpreted as implying any particular order or arrangement among or between various steps or elements except when the order of individual steps or arrangement of elements is explicitly described. EXAMPLES Materials: About 0.55mm thick, about 130 g / m2silver-plated knitted fabric (Shieldex®, Medtex P130) was taken as the conductive fabric. The knit structure incorporated 78% polyamide and 22% elastane, with elastane in both warp and weft directions, imparting two-way stretch properties. The antimicrobial textile composite provided electrical conductivity enabled by silver metallization while maintaining the flexible and stretchable properties of the knitted fabric. This textile is characterized by its high conductive properties, elasticity, and biocompatibility to be used in medical applications. The elastic property enables the application of enough force for the electrodes to meet the skin at the right pressure. The elastomeric polymer (Ecoflex 0030, Smooth-on) employed in the fabrication is characterized by low young’s modulus and elasticity compared to other elastomers which matches the mechanical properties of human skin. Moreover, the sticky nature of the said polymer allows it to adhere to skin firmly yet gently, preventing wrinkling of the textile sensors. EXAMPLE 1: Fabrication of the textile electrode The Ecoflex polymer (elastomeric polymer) was prepared by mixing silicone part A and part B (curing agent) in a 1:1 ratio and cast on a glass plate with a manual applicator. The height of the manual applicator was set to 500 micron as this enabled uniform application of the polymer layer in the textile. The textile was then placed on the silicone layer and cured on a hot plate for about 15 minutes at about 70°C. The silicone entered the porous structure of the textile which helped to adhere the electrodes on the silicone layer after curing. The porous nature of the textile allowed it to partially absorb the thin elastomeric polymer coating. The above-described method yielded a large textile-polymer sensor mat that could be cut into any shape and size. Figure 1 (b) provides a schematic representation of the textile electrode of the present disclosure. For the experiments described below, the textile electrode was cut into patches of dimension of about 4cm x 3cm using a manual fabric rotary blade. EXAMPLE 2: Structural analysis of the textile electrode The textile electrode prepared in Example 1 was subjected to structural analysis by visual inspection right after fabrication and after one or more cycles of washing to understand a) the interaction between the two layers in the obtained textile electrode and b) the durability and / or reusability of the textile electrode. After fabrication, the textile electrode was subjected to structural analysis by using an optical microscope. During the said analysis, it was found that voids formed in the coating due to the rough surface of the textile preserved the inherent conductivity of the textile (Figure 3). In other words, it was seen that the controlled thickness of the elastomeric coating opened small voids in the textile that retain conductivity for capturing ECG signals, enabling continued conductivity. The textile electrode was then subjected to 10 cycles of hand washing to check for its reusability and structural stability. Figure 2b shows the textile electrode after 10 washes where there was no deformation and that the textile electrodes maintained their shape and integrity even after multiple washing cycles (Figure 2b). Figure 2a, on the other hand, for the purposes of comparison, shows that the textile alone, without the elastomeric coating, albeit conductive, cannot be used by itself as electrode due to wrinkling / deformation after one use. It was thus concluded that the insulating elastomeric polymer coating helps retain the textile electrode’s original shape over time, avoiding distortion after repeated use and washing. EXAMPLE 3: Signal Acquisition: Electrical contacts were made to the textile electrode using Thermoplastic polyurethane (TPU) coated silver plated conductive yarns (Shieldex®). These insulated yarns offer advantages over conventional cables, owing to their textile structure and sheathing. Specifically, the yarn is washable and shielded from external interference. Thus, the said yarns provide robust and protected electrical connections to the electrode. Although the polymer layer ensures the electrodes stick to the skin, for additional support, a soft skin adhesive was optionally used to hold the sensors in position. Single lead (two electrodes) ECG measurement was performed with the textile electrodes prepared in Example 1 and compared with the conventional hydrogel electrode (3M red dot electrode). The measurements were all done using an ECG evaluation module from Texas instruments (ADS1292ECG-FE). The electrodes were positioned such that it roughly formed an equilateral triangle called as the Einthoven’s triangle (Figure 4) and connected using the ECG cables for measurement (Biometric Cables). The silicone coated side of the electrode was attached to the skin to form a strong contact. The right electrode (12 cm2area) was placed close to the right arm (RA) and the left electrode (LA) close to the left arm as depicted in the image. The reference electrode was placed on the centre of the chest (LL). Lead 1 was used for recording the measurements and comparison of the signal-to-noise Ratio (SNR). The testing was conducted on one individual participant, following strict adherence to relevant ethical regulations and guidelines. Signal Processing All the signal processing of the signals was done in MATLAB software. The removal of the baseline wander of the ECG signal and filtering were done using a band pass filter. To calculate the noise in the raw ECG signal, the raw ECG signal was subtracted from the filtered signal, The power spectral densities of both the filtered signal and the noise were obtained by applying a fast Fourier transform. Since the area under a power spectral density curve equals the total power, the signal-to-noise ratio (SNR) of the ECG signals was determined by dividing the total power of the denoised ECG signal by the total power of the noise (SNR = Psignal / Pnoise). Characterization Figure 5a illustrates the raw data of the textile electrode compared to the conventional dot tape hydrogel electrode before any washing cycles where the SNR of the textile electrode was about 41.22 dB and the SNR of the hydrogel electrode was about 37.93 dB. It was thus observed that the textile electrode of the present disclosure provides SNR higher than conventional ECG electrodes. To further assess reliability and reusability, textile sensors and conventional electrodes underwent multiple hand washing cycles with mild detergent and air drying. As shown in Fig. 5b, the SNR of the textile electrode decreased by only ~12% (35.92dB) after six washes, with no deterioration in signal quality observed. In contrast, the conventional hydrogel electrodes are single use, so the SNR decreased by 90% (3.92dB) after six washes as the solid conductive gel peels off after one wash which results in a reduced contact to the skin or high contact impedance. The conventional electrodes thus experienced low adhesion and removal of the conductive gel, which led to decreased contact between the electrodes and skin. As a result, the SNR for the conventional electrodes was low. Thus, the textile electrodes of the present disclosure provide durable and robust signal acquisition over repeated use and washing, unlike conventional single-use disposable electrodes. While signal filtering of the conventional hydrogel electrodes (Fig 5d) enhanced the visibility of the ECG R peaks, other signal peaks were not clear and oscillations were found to persist, requiring further processing. Moreover, the conventional hydrogel electrode's adhesive properties degraded with each wash, leading to poor skin adhesion or increased contact impedance. As a result, these conventional hydrogel electrodes were found to be unsuitable for long-term dynamic signal monitoring applications. EXAMPLE 4: Repeatability study between different sets of textile and conventional hydrogel electrodes Three different sets of textile and conventional hydrogel electrodes were tested to compare their signal-to-noise ratio (SNR) over multiple cycles of washing (Figure 6). Initially, the textile electrode had an average SNR of around 34 dB, equivalent to the conventional hydrogel electrode. After 10 wash cycles, the textile electrode SNR remained relatively high at an average of 28 db. In contrast, the hydrogel electrode experienced a significant drop in SNR to about 7 dB (average) after ten washes. The variation in SNR measurements of the textile electrode was also impacted by factors like changes in contact impedance during multiple cycles of tests, variation in electrode size as electrodes are cut manually, and skin condition. Overall, the textile electrode maintained a much higher SNR compared to the conventional hydrogel electrode after repeated washing, suggesting it may be more suitable for applications requiring durability to washing. EXAMPLE 5: Effect of cleaning on sensing ability of the textile electrode To evaluate sensor performance under different cleaning conditions, the electrodes were wiped with isopropyl alcohol (IPA) instead of detergents, as detergents are unsuitable when electronics are involved. Three distinct sets of electrodes were tested, ranging from an uncleaned state to being cleaned ten times with IPA. The observations showed that the conventional hydrogel electrodes became excessively sticky and unusable after the first IPA wipe, making it impossible to record any data beyond a single cleaning cycle. In contrast, the textile electrodes of the present disclosure did not exhibit significant variations in signal-to- noise ratio (SNR) across multiple cleaning cycles with IPA, and maintained an average of 31dB, like when washed with detergent (Figure 7). EXAMPLE 6: Change in conductivity upon varying thickness of the insulating elastomeric polymer coating on the conductive textile The procedure as described in Example 1 was repeated such that thickness of the polymer layer cast on the conductive textile was varied between about 200 µm to about 1000µm. The resistance / conductivity of the said variants of the electrode was measured by using a multimeter. Figure 8 demonstrates the correlation between thickness of the polymer layer and the textile electrode’s conductivity. It was found that conductivity of the textile electrode is significantly impacted by the polymer thickness. At a thickness of 1000μm, the resistance was extremely high, resulting in a complete loss of conductivity. When the thickness was reduced to 800μm, the textile remained conductive, but the resistance was still relatively high such that at certain points, the high resistance caused the textile to behave as if it was non-conductive. Figure 9 shows images of the textile electrode having a polymer layer thickness of about 200μm before and after use. It was seen that not only does the textile deform at thickness below 300μm, but the inherent stickiness that allows it to adhere to skin diminishes due to the extremely thin polymer layer. Accordingly, it was concluded that deviating from polymer layer thickness of 300μm to 600μm yield textile electrodes having unfavorable conductivity and / or structural stability. EXAMPLE 7: Electrodes employing PDMS as non-conductive elastomeric polymer and comparison with the textile electrode of Example 1 The procedure as described in Example 1 was repeated employing PDMS as the non- conductive polymer. The PDMS was prepared using a commercially available kit whereby SYLGARD 184 was mixed with curing agent () at a ratio of about 10:1 as per the manufacturer’s instructions. Although PDMS is also an elastomeric polymer, it possesses a higher Young's modulus compared to Ecoflex, resulting in a change in the mechanical properties of the polymer. The resistance of both the textile electrodes, however, was found to have minimal variation for a polymer thickness of 500µm, as provided below: Resistance of Textile-Ecoflex Electrode = 10.16 ohms Resistance of Textile-PDMS Electrode = 9.42 ohms The above is also depicted in Figures 10 and 11. The SNR of the ECG signal was found to be above >20dB and all the peaks were visible. Accordingly, it was concluded that elastomeric polymers like Silicones (Ecoflex) and PDMS can be employed in the fabrication of the textile electrode of the present disclosure. Additional embodiments and features of the present disclosure will be apparent to one of ordinary skill in art based on the description provided herein. The embodiments herein provide various features and advantageous details thereof in the description. Descriptions of well- known / conventional methods and techniques are omitted so as to not unnecessarily obscure the embodiments herein. The foregoing description fully reveals the general nature of the embodiments herein that others can, by applying current knowledge, readily modify and / or adapt for various applications such specific embodiments without departing from the generic concept, and, therefore, such adaptations and modifications should and are intended to be comprehended within the meaning and range of equivalents of the disclosed embodiments. It is to be understood that the phraseology or terminology employed herein is for the purpose of description and not of limitation. Therefore, while the embodiments in this disclosure have been described in terms of preferred embodiments, those skilled in the art will recognize that the embodiments herein can be practiced with modification within the spirit and scope of the embodiments as described herein, without departing from the principles of the disclosure. Any discussion of documents, acts, materials, devices, articles and the like that has been included in this specification is solely for the purpose of providing a context for the disclosure. It is not to be taken as an admission that any or all of these matters form a part of the prior art base or were common general knowledge in the field relevant to the disclosure as it existed anywhere before the priority date of this application. References: [1] M. Alizadeh Meghrazi et al., “Multichannel ECG recording from waist using textile sensors,” Biomed Eng Online, vol.19, no.1, Jun.2020, doi: 10.1186 / s12938-020-00788-x. [2] C. M. Vidhya, Y. Maithani, and J. P. Singh, “Recent Advances and Challenges in Textile Electrodes for Wearable Biopotential Signal Monitoring: A Comprehensive Review,” Biosensors, vol.13, no.7. Multidisciplinary Digital Publishing Institute (MDPI), Jul.01, 2023. doi: 10.3390 / bios13070679. [3] M. Alizadeh-Meghrazi et al., “Evaluation of dry textile electrodes for long-term electrocardiographic monitoring,” Biomed Eng Online, vol. 20, no. 1, Dec. 2021, doi: 10.1186 / s12938-021-00905-4.

Claims

WE CLAIM:

1. A textile electrode comprising a conductive textile coated with an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns.

2. The textile electrode as claimed in claim 1, wherein the conductive textile comprises a knitted or woven fabric coated with a conductive substance.

3. The textile electrode as claimed in claim 2, wherein the knitted or woven fabric is a porous stretchable fabric.

4. The textile electrode as claimed in any of claims 2 and 3, wherein the knitted or woven fabric comprises polyamide, elastane, cotton and viscose or any combination thereof; and / or wherein the conductive substance is selected from a group comprising metal(s) and conductive carbon-based material(s) or a combination thereof.

5. The textile electrode as claimed in any of claims 1-4, wherein the conductive textile has thickness ranging from about 0.40 mm to about 0.70 mm.

6. The textile electrode as claimed in any of claims 1-5, wherein the insulating elastomeric polymer comprises polymer(s) selected from a group comprising silicone(s) and polydimethylsiloxane (PDMS) or a combination thereof.

7. The textile electrode as claimed in any of claims 1-6, wherein the insulating elastomeric polymer stabilizes shape of the conductive textile and enables stable skin adhesion.

8. The textile electrode as claimed in any of claims 1-7, wherein textile electrode is a reusable textile electrode that retains its shape and conductivity after one or more cycles of washing and / or exposure to cleaning agents.

9. The textile electrode as claimed in claim 8, wherein reduction in SNR after multiple cycles of washing is limited to about 10% to about 25%.

10. A method of fabricating the textile electrode as claimed in any of claims 1-9, comprising: casting the insulating elastomeric polymer on the conductive textile to obtain an insulating elastomeric polymer layer of thickness ranging from about 300 microns to about 600 microns over the conductive textile.

11. The method as claimed in claim 10, wherein the casting is performed at a temperature of about 20°C to about 25°C.

12. The method as claimed in any of claims 10 or 11, wherein the casting step is followed by curing the insulating elastomeric polymer cast on the conductive textile.

13. The method as claimed in claim 12, wherein the curing is performed at a temperature of about 50°C to about 80°C; and / or wherein the curing is performed for about 8 minutes to about 25 minutes.

14. A device for real-time biofeedback comprising the textile electrode as claimed in any of claims 1-9.

15. The device as claimed in claim 14, wherein the device is an electrocardiography (ECG) device for Heart Rate Variability Monitoring (HRV).

16. The device as claimed in any of claims 14 and 15, wherein the device further comprises an additional textile electrode for respiration monitoring to measure Respiratory inductance plethysmography (RIP).

17. The device as claimed in any of claims 14-16, wherein the device is in the form of a skin patch.

18. An ECG monitoring kit comprising the textile electrode as claimed in any of claims 1- 9 or the device as claimed in any of claims 14-17.

19. The ECG monitoring kit as claimed in claim 18, further comprising one or more of an instruction manual, a signal acquisition module, signal processing unit, an ECG evaluation board, an accelerometer, wire(s) and soft skin adhesives for additional support.

20. Use of the textile electrode as claimed in any of claims 1-9 or the device as claimed in any of claims 14-17 in monitoring electrical activity of the heart.

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