A comfortability and signal quality study of conductive weave electrodes in long-term collection of human electrocardiographs

Abstract

Long-term electrocardiogram (ECG) recording can reveal some vital cardiovascular disorders and provide warning of human sudden cerebral or vascular diseases in advance. This requires high-quality ECG skin electrodes. Gel (Ag/AgCl) electrodes were reported to have good signal quality in ECG acquisition, but easily caused human skin irritation or allergy. Consequently, textile electrodes have attracted more attention for long-term ECG acquisition. In this paper, eight woven fabrics with diverse yarns and weft densities were fabricated in plain and honeycomb structures. The fabrics were investigated in terms of comfortability, fabric–skin contact impedance and acquired bio-signal quality. Honeycomb weave electrodes were measured with a high comfort level from subjective and objective views, including pleasant tactile comfort, high visual acceptance, good air permeability and good heat transfer. Weave electrodes made of all conductive filaments in high density had low skin contact impedance and high-quality ECG signals. An increase of compression load on weave electrodes resulted in a decrease of contact impedance with a high signal quality. A conductive honeycomb weave with unit repeat of 6_*6 warps_*wefts presented the highest score of acquired ECG signals of all studied electrodes based on the qualities of the QRS complex, P and T waves, R peak amplitude and variation and signal-to-noise ratio. This study contributes to the future design and fabrication of textile electrodes using honeycomb weave in long-term and real-time collection of human ECGs.

Keywords

comfortability contact impedance electrocardiogram signal textile electrodes

Electrocardiogram (ECG) monitoring of human cardiac activity is non-invasive, and can reveal some vital cardiovascular disorders, for instance, heart rhythm abnormalities known as arrhythmias or dysrhythmias. Some arrhythmias, such as ventricular fibrillation, are emergency conditions that may result in sudden cardiac arrest and death. Others, such as atrial fibrillations, which are subtle to detect, may not be in serious conditions, but may develop over time to fatal consequences, such as blood clotting and stroke.¹ Therefore, a long-term continuous recording of ECGs to properly acquire the heart activity and diagnose possible rhythm disorders in advance is useful to warn of the accumulation of related human chronic diseases, such as Holter monitors in hospital for human ECG recordings of up to 48 hours.²

To acquire ECG signals, a number of conductive biopotential electrodes are usually mounted on multiple positions across the human body. They can interpret the heart activity comprehensively, known as 12-lead ECGs. Conventionally, in hospital, Ag/AgCl electrodes with conductive wet gel are widely adopted. These electrodes, with good gel adhesive to the skin, show lower skin–electrode contact impedance with less motion artifact. However, in some cases there have been reports of human skin irritation and allergic reaction owing to the long-term wearing of gel-based electrodes.³ Meanwhile, the wet electrodes experience a drying process with degradation in long-term wearing,⁴ and the acquired signal thereafter fades gradually. Therefore, the Ag/AgCl wet gel electrodes are not favorable for long-term monitoring applications. Instead, electrodes without gel have been explored,⁵ which appear to be more applicable for long period of cardiac monitoring. Simultaneously, this can provide the patient with desired comfortability when continuously worn as the part of wearable monitoring device.

So far, many “dry” ECG electrodes have been reported, which can be categorized into two types: “contact” and “non-contact” electrodes.⁶ For the “contact” type, the interface between the skin and the electrode is coupling directly, which can be enhanced further in efficiency by some conductive substance, such as sweat or moisture.^7,8 In contrast, the “non-contact” type for acquiring ECG signals is based on capacitive coupling between the electrodes and the skin with a separation layer of dielectric material.⁹ To date, a large number of dry electrodes have been reported, for instance, silicon micromachining for microneedle structural electrodes,¹⁰ three-dimensional (3D) printed electrodes,¹¹ flexible electrodes with thin-film metallic coatings,¹² nanomaterial-doped polymer electrodes¹³ and printed circuit boards (PCBs) for capacitive electrodes.¹⁴

Alternatively, another type of dry electrode has been developed based on conductive fibers or fabrics.¹⁵ Fabric is a material that we frequently use or wear for a long period of time every day. For developing wearable health monitor device, dry electrodes are ideally built using flexible, soft and comfortable fabrics. In the garment manufacture, seamless integration of textile electrodes onto flexible wires, low power and wireless transmission modules and small handheld devices make bio-signal monitoring possible with minimal disturbance to our daily routine life. Moreover, the adopted conductive fixation point,¹⁶ such as the metal snap for connecting electronic components and textile electrodes, has greatly simplified the manufacture process of wearable, smart medical garment for long-term monitoring of human physiological signals. The abundance and variety of clothing accessories, such as snap fasteners, elastic belts and bands, strongly support the integration of miniature electronics on textiles to realize wearable monitoring garments. To date, many types of such garments, such as tight vests, belts, Babaka and shirts worn around the upper chest region, have been explored to sense ECGs with the help of textile-based electrodes embedded in the “garment platform”.^17–20 One main feature of currently developed electronic garments is to offer high tightness of textile electrodes on human skin for the improvement of signal acquisition stability.

The properties of textile electrodes need to be evaluated before mass production, including electrical conductivity, mechanical durability, stability against repeated use, comfort level and flexibility, cost, large-area manufacturability, etc. There have been many fabrication methods to produce textile electrodes with diverse properties, such as weaving, screen printing,^21,22 metallic spinning,²³ conductive filament and fiber blending,²⁴ electroplating,²⁵ etc. In an effort to study textile electrodes, we have recently reported weave-based electrodes and their difference in wearable comfortability and bio-signal acquisition stability.²⁶ The general comparison indicated that honeycomb structural weave electrodes manifested better wearable comfortability and relatively poor quality of acquired ECG signals. However, the inherent rules of how plain and honeycomb weave structures and materials influence the wearing comfortability and ECG acquisition stability are still unknown as yet.

In this paper, eight woven fabrics are fabricated in two structural patterns with diverse yarns and densities. Their properties are investigated in detail, including areal impedance and comfort levels. A tight belt prototype in three-lead mode has been developed to mount the fabricated electrodes as a fully wearable and textile-based garment for long-term ECG monitoring. The belt houses ECG sensing electrodes, miniature electronic circuitry for signal readout and transmission and an energy supply. The belt is inherently elastic, which can result in direct coupling of electrodes to the skin. The acquired ECG signals are then compared under the same conditions. The unique system-level design of the prototype allows us to optimize the front-head material for solving the contradiction of monitoring bio-signal stability and wearable comfortability.

Materials and methods

Woven fabric utilizes warp and weft yarns interlaced perpendicularly with each other with high tightness and structural integrity. This weaving structure is relatively stable in comparison with the easily deformed knitting-loop structure. Thus, the weaving structure should be a better platform for the acquisition of bio-signals due to less motion artifact. Two typical woven fabrics in plain and honeycomb patterns show two-dimensional (2D) and 3D structures.²⁷ Our previous work²⁶ qualitatively investigated the performance of two woven fabrics in ECG acquisition. In the current study, we optimized the internal ECG algorithm and resolution of the display terminal in the ECG acquisition system. Moreover, we quantificationally compare the properties of more woven fabrics systematically, in terms of their wearable properties and bio-signal acquisition functions. This includes subjective and objective evaluations of fabric comfort level, areal/fabric–skin impedances, compression-impedance and bio-signal monitoring characteristics.

Preparation of conductive fabrics for textile electrodes

Conductive nylon (6,6) filaments (70 Denier (abbreviated as 70D, where 1D means the weight (gram) per 9000 meters of filament)) were purchased from Unionweal TH Technology Co. Ltd (Shaoxing, P.R. China). The filament bundles were used to prepare conductive woven fabrics, in which each filament was coated with a thin silver layer, as shown in Figure 1, in which the images were obtained via scanning electronic microscopy (SEM, JEOL Model JSM-6490). The coated conductive silver layer was measured with a thickness range of 100–200 nm, and the layer on the filament is continuous, which gives rise to the electrical resistance of each bundle at 0.02 Ω/cm. To enhance the diversity of fabric samples, other types of yarns were employed into woven fabrics: non-conductive yarns, that is, a pure nylon (6.6) filament bundle without twist as a nonelastic yarn (70D) (Figure 1(b)) and nylon (6,6) filaments wrapped with spandex as elastic yarn (60D).

Figure 1.

Scanning electron microscopy images of (a) a conductive nylon filament with a silver-coated layer, (b) non-conductive nylon filaments without the coated layers and (c) cross-sections of conductive filaments with a central nylon fiber core and coated silver layer on the fiber surface, where (d) shows clearly the coated silver layer and filament core with an evident interface.

Eight woven fabrics with different conductivity and properties were fabricated using the aforementioned materials through a rapier weave loom (SL8900S, an automatic loom for making fabric samples). The fabric specifications are listed in Table 1, in which fabrics (P_SN-26–P_SS-31) were fabricated in a plain weave pattern (

), while fabrics (H_SS18-30 and H_SE18-25) were in a honeycomb weave pattern with 18*18 warps*wefts in the unit repeat (

) and fabrics (H_SS6-32 and H_SE6-30) were in the same pattern with 6*6 warps*wefts in the unit repeat (

). The as-made fabrics were tailored into identical squares with the size of 40×40 mm² to make pillow-shaped electrodes, which contain a thin layer (5 mm in thickness) of sponge between the identical conductive fabrics. The aim of making such shaped electrodes is to ensure the direct coupling of conductive fabric with the human skin, which may avoid interferences for signal collection as occurs in commercial disposable circular electrodes (such as 3 M-brand electrodes) with a diameter of 40 mm.

Table 1.

Specifications of conductive woven fabrics

Fabric	Fabric pattern	Compositions		Warp density	Weft density
Code	(Weave structure)	warp	weft	(yarns/cm)	(yarns/cm)
P_SN-26	Plain ()	S	N	40	26
P_SN-19	Plain ()	S	N	40	19
P_SS-20	Plain ()	S	S	40	20
P_SS-31	Plain ()	S	S	40	31
H_SS18-30	Honeycomb 1818 warpswefts ()	S	S	30	30
H_SE18-25	Honeycomb 1818 warpswefts ()	S	E	30	25
H_SS6-32	Honeycomb 66 warpswefts ()	S	S	32	32
H_SE6-30	Honeycomb 66 warpswefts ()	S	E	32	30

S: conductive filament bundle composite of nylon (6,6) coated with a thin layer of silver, 70D; N: non-conductive filament bundle of nylon (6,6), 70D; E: nylon (6,6) filament wrapping spandex (elastic yarn), 60D; P: plain weave; H: honeycomb weave, SN: combination of conductive S warps and non-conductive N wefts.

Note: the subscript -26 represents the weft density of 26 yarns/cm, and so on. In each fabric pattern, such as , the black spot represents the warp float, while the white spot represents the weft float.

Table 1 lists the specifications of eight conductive fabrics where (P_SN-26 and P_SN-19), (P_SS-20 and P_SS-31), (H_SS18-30 and H_SE18-25) and (H_SS6-32 and H_SE6-30) form four groups. Among them, fabrics (P_SN-26 and P_SN-19) have the same fabric pattern, warp and weft yarns and warp density, but different weft densities. Fabric (P_SN-26) is denser. This situation also applies to the other three groups of fabrics. For each group, the same warp density is due to the limitation of the weaving process once the warping process has been completed and the variation of weft density is because it is easier to change the weft insertion density. The difference of weft densities enables the variation of fabric areal electrical resistance (AER). On the other hand, the first and second fabric groups are different in AER because fabrics (P_SN-26 and P_SN-19) have weft non-conductive yarns causing higher fabric impedance than fabrics (P_SS-20 and P_SS-31). The main difference in the third and fourth groups is displayed through the size of the unit repeat where the third group of fabrics contains 18 warps and 18 wefts (18*18 warp*weft) in the unit repeat while the latter has six warps and six wefts instead. A greater number of warps and wefts in the unit repeat results in thicker honeycomb fabric and larger space in the fabric. In general, the unit repeat size of honeycomb weave influences the physical properties of fabric, contact impedance to skin and the bio-signal quality, which will be investigated further in the following section.

The microstructural morphologies of eight conductive woven fabrics were captured using a Leica M165C microscope, as shown in Figure 2. In general, plain woven fabrics display a relatively flat surface with regular geometry of pores between adjacent warp and weft yarns. The factors of pore shape, size, fabric porosity and thickness determine the fabric properties, such as air permeability, thermal conductivity and wearable comfortability.

Figure 2.

Images of fabric samples, semi-conductive plain woven fabrics with weft density of (a) P_SN-26, 26 yarns/cm and (b) P_SN-19,19 yarns/cm; conductive plain woven fabrics with weft density of (c) P_SS-20, 20 yarns/cm and (d) P_SS-31, 31 yarns/cm; 18*18 honeycomb woven fabric with (e) H_SS18-30, pure conductive warp and weft yarns, (f) H_SE18-25, conductive warp yarns and non-conductive weft elastic yarns; 6*6 honeycomb woven fabric with (g) H_SS6-32, pure conductive warp and weft yarns, (h) H_SE6-30, conductive warp yarns and non-conductive weft elastic yarns.

In contrast, honeycomb weave presents a periodic stereostructure of both surfaces, in which the unit repeat displays inverted pyramidal space. The application of elastic yarns into honeycomb weave strengthens the 3D effect, as shown in the comparison of fabric images in Figure 2 (H_SS18-30 and H_SE18-25). The size of the fabric unit repeat also affects the spatial effect, as shown in Figure 2 (H_SS18-30 and H_SS6-32), where the larger honeycomb unit results in a stronger honeycomb structural effect.²⁷ The stereostructure of the honeycomb weave presents diverse wearable comfortability, such as permeability and moisture transfer,²⁸ as well as improved bio-signal collection quality in comparison with plain weave.²⁶ If we focus on the honeycomb structure of fabric (H_SS18-30), it is found that a part of the warp yarn underneath a weft float is the warp float. Two layers of warp and weft floats form a closed internal space. Owing to the periodical repeats of the inverted pyramidal space on fabric surfaces and the closed internal space inside fabric, the honeycomb weave has been reported with high wearable comfort level, for instance, good sound and moisture absorption, good insulation and a quick dry rate.²⁸ In this paper, the four fabricated honeycomb weaves display different patterns, electrical resistance and bio-signal acquisition ability due to the different contents of conductive filaments in the fabric unit repeat.

Measurement of contact impedance of skin–electrode

Measuring the interface impedance of the skin and the electrode is an important way to evaluate the performance of as-made weave electrodes in the acquisition of ECG signals. Referring to our previous work in determining the impedance,²⁶ a simpler approach was adopted, as shown in Figure 3(a). A digital multimeter was employed to measure the impedance from one electrode passing over two fabric–skin interfaces and a skin area. The reference electrode using Ag/AgCl wet gel is from Solar Co.; the “Electrode” in Figure 3(a) can be replaced with fabricated weave electrodes. The skin–electrode impedances can be calculated as²⁶

Z_{pq} = Z_{q 1} + Z_{e 1} + Z_{e 2}

(1)

Z_{qr} = Z_{q 2} + Z_{e 2} + Z_{y}

(2)

where

Z_{q 1}

and

Z_{q 2}

denote the body impedance between the points p, q and q, r, respectively. These values differ from person to person.

Z_{e 1}

and

Z_{e 2}

denote the interface impedances between the skin and the reference electrode, respectively. Z_y is the interface impedance between the skin and the weave electrode.

Z_{pq}

and

Z_{qr}

represent the impedances between measured electrodes. In order to compare the difference of commercial electrodes and self-developed weave electrodes, Z_y to Reference 2 is located at the same distance as Reference 1 to Reference 2, that is,

L_{1} = L_{2}

.²⁹ Because the same reference electrode was used in this study, Equations (1) and (2) are simplified as

Z_{e 1} = Z_{e 2}, Z_{q 1} = Z_{q 2}

(3)

Z_{y} = (Z_{qr} - Z_{q 2}) / 2

(4)

Figure 3.

Contact impedance measurement of electrodes and skin, (a) principle of measurement, where L₁ represents the distance between References 1 and 2, while L₂ represents the distance between Reference 2 and the electrode; and (b) real setup for measuring the electrode–skin impedance.

The electrode impedance Z_y was calculated using Equation (4), using only the measured values of $Z_{qr}$ and $Z_{q 1}$ . Here, the contact impedance (Z_y) between the fabric and the skin is affected by the fabric pattern, conductive materials and geometric parameters, as well as the compression force on the fabric. Therefore, the compression force should be constant in later bio-signal acquisition. In this study, the AERs of as-made weave electrodes were measured with the increase of compressions and the values were compared with the contact impedance to find the rules of ECG acquisition.

Physical properties of weave electrodes

Long-term wearing of wet electrodes may result in human skin irritation or allergic reaction. Thus, high comfortability is required for people to accept the alternative electrodes. The internal geometrical structure, yarn density and materials of electrodes decide the physical properties. The physical properties, such as air/moisture permeability, thermal conductivity, tactile comfort, etc., determine the comfort levels of the electrodes. In general, the comfortability report of an electrode should include subjective and objective evaluations. A number of subjective comfort indexes, such as tactile comfort and the style, require volunteers to wear electrodes to feel them and mark the comfort levels. For example, the tactile comfort of the weave electrode is at a high level when the weave has the desirable microstructure and warp/weft density. The discomfort is mainly because of tickle, scratchiness, itch, prickle, fiber shedding and other factors that can irritate the wearer’s skin during the long-term wearing.

This study involves eight fabricated weave electrodes with eight unique comfortabilities. In order to obtain the subjective comfort level of electrodes, each sample was attached on the self-developed wearable system. The wearable system offered the electrodes a tight contact to chest skin under a constant compression pressure (e.g. 50 Pa). Twenty volunteers participated in the evaluation, and each participant was asked to complete a feeling questionnaire on comfort levels after one day of wearing. The questionnaire contains questions, for example, how is your feeling of the fabric tactile comfort and visual acceptance? The answers were scored with five grades, that is, perfect (5), good (4), satisfactory (3), not good (2), poor (1). The subjective feeling for each weave electrode was scored with the average value with the standard deviation that was provided by all participants.

The objective evaluation of weave electrodes, such as air permeability (resistance) and thermal conductivity, to some extent can give us the comfortability of the developed electrodes objectively. For example, high air permeability of electrodes can transfer sweat and moisture properly, which results in good comfortability for people. The weave-based electrodes are measured for air permeability and thermal conductivity using KES set equipment,²⁸ that is, the KES-F8-AP1 tester for measuring air resistance and the KES-F Thermo Labo II for measuring thermal conductivity. For the KES-F8-AP1 tester, the speed of the piston inside the tester was set as 2 cm/s and the cylinder section area was 4π cm²; therefore, the air flow rate was 8π cm³/s. The air resistance (R) is defined as the pressure difference divided by the air flow rate per unit area, which has a unit of KPa s/m. Regarding the KES-F Thermo Labo II, the power loss through the fabric sample from the BT-Box (watt) to the Water Box was measured. The temperature of the BT-Box and Guard Water Box were set to 30℃ and 20℃, respectively. Based on the power consumption of the test plate heater, the amount of heat passing through the sample (in watts per square meter) was recorded. The thermal conductivity value (λ), having a unit of W/m K, for the honeycomb and plain woven fabrics can be calculated by Equation (5)

λ = w \times \frac{s}{A \times Δ T}

(5)

where w is the heat flow lost, s is the sample thickness, A is the heat contact area of “BT-Box” and ΔT is the temperature difference of the sample. Each final objective value was scored as the average value obtained from five measurements for each parameter.

Establishment of a wearable ECG monitoring system

Upon verifying the performance of ECG measurement using self-developed weave electrodes, a complete, wearable prototype of a belt with integrated electronics was developed. The wearable belt is a platform for mounting textile electrodes, as shown in the in-kind picture in Figure 4. The fine wires that connected the three leads and central processor are embedded inside the belt. A small handheld device contains a lithium battery and an electronic module that integrates the chip, PCB, signal transmission module, etc. Here, low-cost, commercially available discrete components were used; the ECG acquisition circuitry was custom-designed. The acquired ECG signal can be transferred to a self-developed terminal displayer app on a smart phone. The pillow-shaped textile electrodes with front and reverse sides and a reference electrode (Ag/AgCl) are also shown in Figure 4. The distribution of the three electrodes corresponding to the easy position of ECG acquisition on the human body is illustrated on the left-hand side of Figure 4. The obtained one-lead is approximately oriented along the cardiac axis. This type of lead is used to produce a high-quality ECG signal because the typical orientation of the cardiac axis is well-known clinically. In addition, this one-lead of standard placement has an intrinsic advantage in conveying useful information for inspection. In this study, an Ag/AgCl electrode was used as a ground reference, and it was placed side by side with weave-based electrodes for recording ECG signals.

Figure 4.

A protocol for establishing an electrocardiogram (ECG) monitoring system with electrode placement, real wearing image, in-kind wearing system, electrode samples and a standard ECG signal containing P-QRS-T waves.

For the electrode placement in detail, as shown in Figure 4, two identical textile electrodes are positioned on the chest surface (positions #l and #2); the third reference electrode (#3) is 10 cm away from the left top electrode #1. For example, when as-made fabric (P_SN-26) was tested for ECG signal acquisition, two textile electrodes made of fabric (P_SN-26) were placed in the positions of #l and #2. Position #3 was attached to an Ag/AgCl gel electrode. The ECG measurement using other fabricated fabrics was identical to that of fabric (P_SN-26). This arrangement of electrodes imitates the usage of electrodes with Holter in hospital, because the close distance to the human heart enables one to acquire relatively clean ECG signals with higher R-peak amplitude and QRS complex waves. The pure conductive fabric surface of the pillow-shaped textile electrode contacts the human skin, and the other side with a metal snap connects to the belt wire. The acquired ECG signals were visualized through a smart phone, and the related ECG data were stored in a base station.

Evaluation of acquired ECG signals

One right-handed volunteer who had undergone heart surgery once was employed to test the textile electrodes for ECG acquisition. The measurement conditions were set to be standard that the ambient temperature was 25℃ and relative humidity was 65%; the ECG monitoring was carried out at night for half an hour after refraining from smoking and coffee utilization for six hours before this measurement. The ECG signals were captured in four sections after the wearable system was put on the volunteer: 10 s, 1 min, 5 min and 30 min. The acquired ECGs were compared. In fact, the ECG signal quality was influenced by many factors. Here, we only consider the effect of weave electrodes on ECG signals. On the other hand, to evaluate the signal quality, a method has been developed with variable weighting coefficients to objectively assess the ability of textile electrodes in the acquisition of ECG signals.

The objective evaluation method considers the unit repeat of the ECG signal, in which the QRS wave cluster, P wave and T wave are the main characteristics for describing the heart cyclic action. As shown in Figure 4, a standard analog ECG signal is given clearly that shows the position and configuration of the five waves. The P wave is to the left of the QRS waves and the T wave is to the right of the waves; the P wave is slightly difficult to recognize in comparison with the T wave, but it can reflect many heart rhythm abnormalities in advance. In QRS waves, the direction R wave is the same as the P and T waves, and Q and S waves are in the opposite direction. Normally, the S wave is easier to detect compared with the Q wave. Many heart features, such as heart rate, can be reflected through the R wave characteristics. To evaluate the signal quality, a marking score of 2 is given for the acquired ECG signal with clearly visible and no noise (low signal-to-noise ratio); a score of 0 means the baseline was very noisy (high signal-to-noise ratio). In addition, 2 points are scored when the period from dressing on the belt to the acquired stable ECG signal is less than 10 s; 1 point is obtained for the stabilization time between 11 and 60 s; otherwise, 0 score is obtained.

During the ECG measurement, the belt was dressed in the same position the entire time. In this case, the R-peak amplitude indicates the signal strength obtained from the weave-based electrodes. Here, we suppose that the contact impedance of fabric to skin decides the R-peak amplitude. A 2 point score is obtained when the amplitude is over 0.30 mV, and 0 points are given when the amplitude is less than 0.10 mV. Moreover, the variation of R-spike amplitude indicates the stability of the ECG signals, and the small variation indicates a stable ECG signal. A 2 point score is obtained when the variation is less than 0.04 mV, and 0 points are given when the variation is over 0.08 mV.

Table 2 lists the weighting coefficients for the final evaluation of textile electrodes in ECG monitoring and analysis, including the QRS wave cluster (30%), P wave and T wave (20%), R wave amplitude (20%), stabilization time (20%) and R-peak amplitude variation (10%).²⁶

Table 2.

Weighting coefficients for the final electrocardiogram signal evaluation²⁶

Particulars	QRS-complex	P-wave	T-wave	R wave amp	Stabilization time	R-peak amp variation
Weight coefficient	30%	10%	10%	20%	20%	10%

The weighting coefficient was suggested from a senior doctor who has experience in the diagnosis of cardiovascular diseases through acquired ECG patterns. The coefficients also indicate the importance of each item in judging the related diseases. For example, 30% of weight coefficient of the QRS-complex means that QRS waves contain around 30% of the information relating to cardiovascular diseases. These coefficients in the ultimate scoring decide the final comprehensive assessment of the developed weave electrodes. Therefore, they are important in the later design of textile electrodes for comparison of the obtained ECG signals and standard signals from Holter, and thereafter the guiding practical diagnosis of cardiovascular diseases.

Results and discussion

Comparison of physical properties of weave electrodes

The surface morphologies of fabricated woven fabrics are shown in Figure 2, in which the honeycomb weave (18*18 warps*wefts) manifests a 3D structure with a rough surface. The related fabric thickness was measured using a standard tester (6 g/inch² pressure).

Correspondingly, in Figure 5(a) the measured thickness shows that the honeycomb weave is much thicker than the plain weave. Among honeycomb weaves, the involvement of elastic yarns in fabric enhances the fabric thickness significantly; fabric (H_SE18-25) shows the thickest fabric of all. The non-conductive yarns show larger diameters than conductive filaments, which causes fabric (H_SE6-30) to be thicker than fabric (H_SS18-30), and fabrics (P_SN-26 and P_SN-19) are thicker than fabrics (P_SS-20 and P_SS-31). The thicker fabric usually shows the higher warmth retention, and with internal close space reported,²⁷ honeycomb woven fabric indicates a relatively higher level of comfort.

Figure 5.

Measured physical properties of (a) thickness, (b) thermal conductivity, (c) air resistance and (d) areal electrical resistance of eight fabricated woven fabrics. Note: the abscissa in each histogram represents the eight woven fabrics, namely, (a) P_SN-26; (b) P_SN-19; (c) P_SS-20; (d) P_SS-31; (e) H_SS18-30; (f) H_SE18-25; (g) H_SS6-32; (h) H_SE6-30.

Evidently reflected by the test results of thermal conductivity shown in Figure 5(b), plain woven fabrics show higher thermal transfer ability than honeycomb woven fabrics, because the conductive filament was coated with a silver layer that provides inherently high thermal conductivity. The thin fabric thickness of plain woven fabric is also a reason for its good thermal transfer; in contrast, the concave surface of honeycomb woven fabric, especially fabric (H_SE18-25), indicates the best warmth retention because the inside spatial structure can store a large amount of still air for low heat conduction and convection. This honeycomb weave is useful for wearers in some extreme weather conditions.

Air resistance shows the difficulty of air transfer through a porous material. It depends on the porosity of the fabric and the pore structure. A larger value of air resistance means more difficult air or moisture transfer through the fabric. In Figure 5(c), fabrics (P_SN-26), (P_SN-19) and (H_SE6-30) show relatively higher air resistance value, especially fabric (P_SN-26), indicating the planar structure of plain woven fabric with high tightness. This tight structure may not benefit the comfortability of fabric because the tight interlacing structure obstructs the exchange of air or moisture at the microenvironment of the fabric and skin. In other words, for the objective assessment of as-made conductive fabrics for textile electrodes, honeycomb conductive weaves seem better candidates for long-term ECG monitoring.

Table 3 lists the subjective scores in average values with standard deviations for the eight fabricated weave electrodes. Each score was obtained from 20 volunteers of their subjective feelings on the attached electrodes. For example, regarding the tactile comfort, each fabric was attached to the skin via a tight belt for six hours. In fact, no volunteer felt uncomfortable during the practice. This is the big difference in woven fabric with wet gel electrodes. However, the scores of honeycomb woven fabrics in wearing feel are generally higher than those of plain woven fabrics. This may be because there is more interwoven space in the honeycomb weave with higher air permeability and lower heat transfer. The scores of visual acceptance in Table 3 indicate that most people have no repugnance for the developed textile electrodes. Visually, they accepted the textile electrodes for their long-term ECG monitoring.

Table 3.

Subjective evaluation of developed textile electrodes

Textile electrodes	P_SN-26	P_SN-19	P_SS-20	P_SS-31	H_SS18-30	H_SE18-25	H_SS6-32	H_SE6-30
Tactile comfort (?/5)	3.5 ± 0.3	3.5 ± 0.4	4 ± 0.2	4 ± 0.2	4.5 ± 0.3	4.5 ± 0.3	4.5 ± 0.2	4 ± 0.2
Visual acceptance (?/5)	4 ± 0.3	4 ± 0.3	3.5 ± 0.3	3.5 ± 0.3	4.5 ± 0.3	4.5 ± 0.3	4.5 ± 0.2	4.5 ± 0.2

Note. The symbol ‘?' means the score obtained from the full score of 5.

Impedance performance of weave electrodes

Figure 5(d) shows the measured AER results of the eight fabricated woven fabrics. The conductive fabrics, (P_SS-20), (P_SS-31), (H_SS18-30) and (H_SS6-32), show lower AER values in comparison with the fabrics involved with non-conductive filament bundles, such as fabrics (P_SN-26) and (H_SE18-25). Most fabrics manifest AER less than 10³ Ω/m², with the same order of magnitude of wet gel electrodes. For each fabric group, the AER is decreased significantly with the increased density of conductive yarns. For instance, fabric (P_SS-31) has a weft density of 30 wefts/cm, in comparison with fabric (P_SS-20) with a density of 20 wefts/cm: the measured result showed that the AER of fabric (P_SS-31) is almost one fifth of the value of fabric (P_SS-20). With the increased density of non-conductive filament bundles, the AER has also been improved significantly, as shown in the measured values of fabric P_SN-26 and P_SN-19, which increase from 2564 to 25,420 Ω/m².

In fact, Figure 5(d) reflects the static values of AER of as-made woven fabrics for each fabric under a certain areal compressed pressure. When the as-made conductive fabrics were compressed under a series of compression pressures, the fabric AER would reveal a kind of dynamic value correspondingly. The test model is shown as the inset illustrated scheme in Figure 6(a), where the conductive woven fabric was held up and compressed by a pair of parallel metal plates. All fabric samples were tailored at the same size. The compressed loadings were exerted on one plate, and the corresponding resistances were recorded through a digital multimeter. When the compression loading is close to zero, the measured resistance divided by the fabric test area is the fabric static AER. An increased compression loading then gives the measured relationship of dynamic electrical resistances and exerted pressures, as shown in Figure 6(a).

Figure 6.

The measured relationship of electrical resistance and exerted pressure for the eight woven fabrics, namely, (a) P_SN-26; (b) P_SN-19; (c) P_SS-20; (d) P_SS-31; (e) H_SS18-30; (f) H_SE18-25; (g) H_SS6-32; (h) H_SE6-30, respectively. (a) Fabrics (c and d) in which the inset image shows the measure principle. (b) Fabrics a, b and e–h.

The measured results in Figure 6 all displayed a decreasing trend with a rapid decline at the beginning of compression, and a slow decrease afterwards. This is especially observed in fabrics (P_SN-26), (P_SN-19) and (H_SE18-25), indicating that high compression can lead to high electrical conductivity. The reason for this is ascribed to the enlarging contacted area of conductive filaments at the interlaced position under increasing compression pressure. For fabrics (P_SS-20), (P_SS-31), (H_SS18-30) and (H_SS6-32), the declining trend is not the same for fabrics (P_SN-26) and (P_SN-19); because they were fabricated with conductive warp and weft filaments, the conductivity is not changed as much as fabrics with non-conductive yarns.

In this article, wet gel Ag/AgCl electrodes as reference electrodes were used to measure the contact impedance as mentioned in the test principle in Figure 3. The test results are listed in Table 4, in which the impedance between two reference electrodes (R+R) on the forearm shows 4.5 Megohm (MΩ) per 10 cm. The other eight values were obtained using a reference electrode and an as-made conductive textile electrode. As for the tested values, it is noted that the contact impedance is increased with the increase of fabric electrical resistance. Here, fabrics (P_SN-26), (P_SN-19), (P_SS-20) and (H_SE18-25) all show high contact impedance due to the non-conductive filament bundles in fabrics (P_SN-26), (P_SN-19) and (H_SE18-25), and the loose structure of fabric (P_SS-20). The smaller AER of conductive fabric results in smaller contact impedance when comparing the values of R+P_SS-20 and R+P_SS-31. The closest value of contact impedance to the R+R value is fabric (H_SS6-32), due to the dense structure of conductive filaments and floats parallel to the skin. This value is slightly larger than the R+R value, indicating the advantage of wet gel in reducing the interface impedance compared with dry conductive materials. This is because the wet gel shows good shape fitting with human rough skin surface. Compared with the AER of conductive fabrics, the contact impedance of the fabric to skin shows many orders of magnitude higher, indicating that the main factor influencing ECG signal monitoring should be the contact impedance of fabric to skin. The subsequent study should focus on the reduction of interface impedance of dry electrodes on the skin without causing a skin allergenic reaction.

Table 4.

The measured impedance between fabric with human skin (R means Ag/AgCl gel electrodes)

Fabric group	R + P_SN-26	R + P_SN-19	R + P_SS-20	R + P_SS-31	R + H_SS18-30	R + H_SE18-25	R + H_SS6-32	R + H_SE6-30	R + R
Impedance (MΩ)	12.7	10.9	8.1	5.7	6.3	11.9	4.8	9.8	4.5

ECG signal evaluation acquired from textile electrodes

The acquired ECG signal was sampled at 250 Hz by the built-in, 10-bit analog-to-digital converter of a commercial microcontroller board (Shenzhen Benefm Co, China), and transmitted through a Bluetooth module (BlueSMiRF silver, SparkFun, Boulder, Co, USA) to a mobile smart phone. The Bluetooth transmission, sampling rate, communication protocol and signal filtration were all set as constant during the monitoring. In this study, the textile electrode was the only variation factor, as shown in the four-moment screenshots of ECG signals using each type of weave electrodes in Figure 7, which gathers 32 moments of ECG signals divided into eight groups related to the eight fabricated fabrics.

Figure 7.

Electrocardiogram (ECG) signals acquired from conductive textile electrodes made of fabrics (a) P_SN-26; (b) P_SN-19; (c) P_SS-20; (d) P_SS-31; (e) H_SS18-30; (f) H_SE18-25; (g) H_SS6-32; (h) H_SE6-30, in each group of ECG signals composed of four ECG moments from signal unstable to stable status (in each ECG grip map, represnts 1 mV in longitudinal coordinate).

In general, weave electrodes made of fabrics (P_SS-31), (H_SS18-30) and (H_SS6-32) give relatively ideal ECG signals, especially fabric (P_SS-31), in terms of the acquired signal quality of the QRS wave complex, T and P waves, R wave amplitude, signal-to-noise ratio, signal stabilization time and R-peak amplitude variation. The ECG signal is neater, clearer and easier to recognize with the increase of conductive filaments per unit area, as shown in the ECG signal comparison from fabrics (P_SS-20) and (P_SS-31) in Figure 7(b).

A standard analog ECG signal is given in Figure 4, which was a model for comparing the measured ECG signal qualities. In detail of score ranking for each investigated factor, the senior doctor suggested that the signal quality ranking of the QRS wave complex and signal-to-noise ratio from high to low in Figure 7 was as follows: (d), (g), (e), (c) (score = 2); (f), (b) (score = 1); (h), (a) (score = 0). For P and T waves, the ranking was as follows: (d), (g), (c) (score = 2); (e), (f) (score = 1); (h), (b), (a) (score = 0). In respect of R wave amplitude, the ranking was as follows: (a), (b), (g) (score = 2); (e), (h), (f) (score = 1); (c), (d) (score = 0). For signal stabilization time from short to long term, the ranking was as follows: (d), (g), (e) (score = 2); (h), (c), (b) (score = 1); (f), (a) (score = 0). As for the R-peak amplitude variation, the ranking was as follows: (c), (f), (h) (score = 2); (d), (g), (a) (score = 1); (e), (b) (score = 0). According to the weighting coefficients for the final ECG signal evaluation in Table 2, a calculation for the above scores gives a comprehensive ranking of the textile electrodes in acquiring ECG signal ability, as follows: (g) H_SS6-32, 1.9; (d) P_SS-31, 1.6; (c) P_SS-20, 1.4; (e) H_SS18-30, 1.4; (b) P_SN-19, 0.9; (f) H_SE18-25, 0.9; (h) H_SE6-30, 0.6; (a) P_SN-26, 0.5. Here, the electrodes made of conductive fabrics, (H_SS6-32), (P_SS-31), (P_SS-20) and (H_SS18-30), show the comprehensive scoring of more than 1.2,³⁰ indicating that these conductive fabrics seem to be suitable for bio-signal monitoring in terms of signal visual quality and easy recognition of some cardiovascular disorders. The characteristics of such fabrics were fabricated with all conductive yarns. The denser fabric structure gives a higher quality of acquired ECG signals. Moreover, the score ranking for ECG acquisition is almost consistent with the contact impedance measured results, in which the smaller contact impedance results in better ECG signal quality. This may be ascribed to the interface coupling such that the higher contact impedance has an easier motion artifact with a less effective contact area of conductive filaments with skin.

Considering the measured comfort level of the eight textile electrodes (the Comparison of physical properties of weave electrodes section), the honeycomb structural woven fabric shows higher comfortability. Moreover, fabric (H_SS6-32) showed the top ranking in ECG acquisition; thus, it can be concluded that the honeycomb structural weave with small size of the unit repeat or a small number of conductive filaments (6*6 warps*wefts) in the unit repeat is a more suitable candidate for textile electrodes in bio-signal monitoring.

Conclusions

Owing to the characteristics of flexible, breathable and biocompatible features of textiles in comparison with wet conductive gels, textile electrodes have been widely used in the development of wearable electronics for bio-signal acquisition. For some human chronic cardiovascular and cerebrovascular diseases, long-term monitoring of ECG signals requires a high comfort level of wearable electronics, which seems fitter for textile electronics. Our practical experience tells us that knitted structural electrodes easily result in the motion artifact of signals due to the easy loop deformation of electrode wearing. Therefore, in this paper, eight fabrics in plain and honeycomb weaves were fabricated into four groups for the quality comparison of acquired ECG signals.

Three aspects of experiments were performed in terms of comfortability, contact impedance and ECG acquisition performance. Regarding comfortability, the honeycomb weave shows higher air permeability and thermal resistance, better tactile comfort and visual acceptance compared with the plain weave. In respect of impedance, the fabrics made of all conductive filaments show low AER, as well as corresponding contact impedance. The contact impedance influences the ECG monitoring such that a lower value of contact impedance causes a better ECG signal quality. In ECG signal evaluation, a comprehensive scoring of the eight weave electrodes gives a final ranking from the best ECG signal to the poorest signal acquisition, based on the set weighting coefficients. Fabrics with all conductive filaments produce the best ECG signal quality, while non-conductive filaments involved in fabrics would bring in a high signal-to-noise ratio and motion artifact. This is an indication that the spatial structure of the honeycomb weave with a high comfort level should be used in the development of textile electrodes after comparison of subjective and objective evaluations. Moreover, the honeycomb weave as electrodes with a small number of warps and wefts in the unit repeat using all conductive yarns can also acquire high-quality ECG signals. Thus, a conclusion was given that such a honeycomb weave can be used for long-term ECG monitoring.

Footnotes

Authors' note

Yanjia Gu is also affiliated with Shenzhen Digital Life Institute, Shenzhen, P.R. China.

Declaration of conflicting interests

The authors declared no potential conflicts of interest with respect to the research, authorship and/or publication of this article.

Funding

The authors disclosed receipt of the following financial support for the research, authorship, and/or publication of this article: This work was supported by “The open fund project of Jiangsu provincial scientific research platform” (Grant Number YGKF-201712) and the National Undergraduate Entrepreneurship Program (Grant Number 1065210232174890) and the Undergraduate Innovation Practice Program (Grant Number1065210232184080).

References

Liu

Tang

Luo

et al.

Wearable carbon nanotubes-based polymer electrodes for ambulatory electrocardiographic measurements. Sens Actuat A Phys 2017; 265: 79–85.

Fung

Jarvelin

Doshi

et al.

Electrocardiographic patch devices and contemporary wireless cardiac monitoring. Frontier Physiol 2015; 6: 149–149.

Peng

Liu

Tian

et al.

A novel passive electrodes based on porous Ti for EEG recording. Sens Actuat B Chem 2016; 226: 349–356.

Marozas

Petrenas

Daukantas

et al.

A comparison of conductive textile-based and silver/silver chloride gel electrodes in exercise electrocardiogram recordings. J Electrocardiol 2011; 44: 189–194.

Yao

Zhu

. Nanomaterial-enabled dry electrodes for electrophysiological sensing: a review. JOM 2016; 68: 1145–1155.

Sun

. Capacitive biopotential measurement for electrophysiological signal acquisition: a review. IEEE Sens J 2016; 16: 2832–2853.

Mcadams E. Biomedical electrodes for biopotential monitoring and electrostimulation. In: Yoo HJ and van Hoof C (eds) Bio-medical CMOS ICs. 1st ed. New York: Springer, 2011; pp.31–124.

Zhang

Tao

. Textile-structured electrodes for electrocardiogram. Text Progr 2008; 40: 183–213.

Pallas-Areny

. Comments on “Capacitive biopotential measurement for electrophysiological signal acquisition: A review”. IEEE Sens J 2017; 17: 2607–2609.

10.

Tay

FEH

Guo

et al.

A microfabricated electrodes with hollow microneedles for ECG measurement. Sens Actuat A Phys 2009; 151: 17–22.

11.

Salvo

Raedt

Carrette

et al.

A 3D printed dry electrodes for ECG/EEG recording. Sens Actuat A Phys 2012; 174: 96–102.

12.

Weder

Hegemann

Amberg

et al.

Embroidered electrodes with silver/titanium coating for long-term ECG monitoring. Sensors 2015; 15: 1750–1759.

13.

Asyraf

Anwar

Sheng

et al.

Recent development of nanomaterial- doped conductive polymers. JOM 2017; 69: 2515–2523.

14.

Jain

Hole

Deshmukh

et al.

Dynamic capacitive sensing of droplet parameters in a low-cost open EWOD system. Sens Actuat A Phys 2017; 263: 224–233.

15.

Rai

Shyamkumar

et al.

Nano- bio textile sensors with mobile wireless platform for wearable health monitoring of neurological and cardiovascular disorders. J Electrochem Soc 2014; 161: B3116–3150.

16.

Choe WC and Ruzumna P. Electrocardiograph monitoring device and connector. Patent 8738112, USA, 2014.

17.

Paul

Torah

Beeby

et al.

A printed, dry electrodes Frank configuration vest for ambulatory vectrocardiographic monitoring. Smart Mater Struct 2017; 26: 025029–025029.

18.

Opio

Kellett

. An observational study of the quality of ECGs recorded by inexperienced staff in a resource-poor African hospital using a reusable ECG belt linked to an internet ECG device. Eur J Intern Med 2017l; 42: E17–E18.

19.

Balsam

Lodzinski

Tyminska

et al.

Study design and rationale for biomedical shirt-based electrocardiography monitoring in relevant clinical situations: ECG-shirt study. Cardiol J 2018; 25: 52–59–52–59.

20.

Boehm

Neu

et al.

A novel 12-lead ECG T-shirt with active electrodes. Electronics 2016; 5: 75–75.

21.

Paul

Torah

Beeby

et al.

The development of screen printed conductive networks on textiles for biopotential monitoring applications. Sens Actuat A Phys 2014; 206: 35–41.

22.

Merritt

Nagle

Grant

. Fabric-based active electrodes design and fabrication for health monitoring clothing. IEEE Trans Inform Technol Biomed 2009; 13: 274–280.

23.

Liu

Pan

et al.

A flexible sensing device based on a PVDF/MWCNT composite nanofiber array with an interdigital electrodes. Sens Actuat A Phys 2014; 211: 78–88.

24.

Tsukada

Nakashima

Torimitsu

. Conductive polymer combined silk fiber bundle for bioelectrical signal recording. PLoS One 2012; 7: e33689–e33689.

25.

Yoon

Kim

et al.

Nanofiber web textile dry electrodes for long-term biopotential recording. IEEE Trans Biomed Circuit Syst 2013; 7: 204–211.

26.

Xiao

Pirbhulal

Dong

et al.

Performance evaluation of plain weave and honeycomb weave electrodes for human ECG monitoring. J Sens 2017; 5: 1–13–1–13.

27.

Xiao

Hua

et al.

Geometrical modeling of honeycomb woven fabric architecture. Text Res J 2015; 85: 1651–1665.

28.

Xiao

Hua

Wang

et al.

Transfer and mechanical behavior of three-dimensional honeycomb fabric. Text Res J 2014; 85: 1281–1292.

29.

Taji

Shirmohammadi

Groza

. Impact of skin–electrodes interface on ECG measurements using conductive textile electrodes. IEEE Trans Instrum Meas 2014; 63: 1412–1422.

30.

Dai

Xiao

Chen

et al.

A low-power and miniaturized electrocardiograph data collection system with smart textile electrodes for monitoring of cardiac function. Aust Phys Eng Sci Med 2016; 39: 1029–1040.