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Open access peer-reviewed chapter

Factors Affecting Wearable Electrode Performance and Development of Biomimetic Skin Phantom

Written By

Krittika Goyal and Steven W. Day

Submitted: 02 December 2022 Reviewed: 23 March 2023 Published: 18 April 2023

DOI: 10.5772/intechopen.111429

New Advances in Biosensing IntechOpen
New Advances in Biosensing Edited by Selcan Karakuş

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New Advances in Biosensing

Edited by Selcan Karakuş

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Abstract

In-home physiological monitoring devices enable the monitoring of vital health parameters and can facilitate health recovery. The current state of the art is inclined towards non-invasive technologies such as wearable mobile devices and patch-based sensors. In this chapter, we provide an overview of progress made in the field of dry electrodes for biopotential acquisition, based on their mechanistic principles, materials, testing methods, and effectiveness in a real-world setting. Important parameters affecting the dry electrode performance such as the area, material, applied pressure and skin hydration are discussed. Traditionally, the development and testing of these wearable electrodes are conducted empirically, in vivo on human skin. However, due to the inter- and intra-subject variability in human skin properties, reliability, repeatability, and the efficacy of the device under investigation cannot be evaluated. Thus a review is presented about the skin phantoms used to simulate the electrical properties of the skin, which has the potential to serve as a robust method to test the functionality of current and future electrodes. This retrospective overview provides researchers with an understanding of the mechanistic principle of biopotential electrodes and the crucial factors that affect electrode performance, thus facilitating wearable electrode development.

Keywords

  • dry electrodes
  • skin phantoms
  • biopotential signals
  • wearable devices
  • non-invasive
  • skin-electrode impedance

1. Introduction

In-home physiological monitoring, also known as remote patient monitoring, enables long-term tracking of patients’ health from the comfort of their homes without the need to visit a clinic. It has the potential to facilitate a healthcare transformation from reactive to proactive preventive care. In-home physiological monitoring of biopotential (voltage) signals such as an Electroencephalogram (EEG), Electrocardiogram (ECG), and Electromyogram (EMG) enables tracking daily changes in a patient’s health [1]. The current state of the art for in-home monitoring is inclined towards non-invasive technologies such as wearable mobile devices and patch-based sensors (Figure 1). In both technologies a biopotential electrode picks up voltage signals from the surface of the body and acts as an interface between the tissue and the electronics.

Figure 1.

Biopotential signals acquisition such as electroencephalogram (EEG), electrocardiogram (ECG), and electromyogram (EMG) using wearable electrodes on the skin.

This chapter is focused on the mechanism of dry electrodes, but we include here a brief description of the most common type of biopotential electrode, the wet electrode. A wet electrode is an electrochemical electrode and consists of two major parts: the metal plate and the electrolyte. An example of a wet electrode is an ECG disposable electrode commonly used in clinical settings. It consists of an electrochemically active gel (this acts as an electrolyte and contains conductive ions), which is in direct contact with both the skin and the metallic part of the electrode, the inside of which is coated with Ag/AgCl and the outside is connected to a metal snap. The whole assembly is mounted on a flexible foam pad and contains adhesive for ease of attachment to the body. Due to the adhesive, wet electrodes stay in contact with the skin while the person is moving and can compensate for the motion artifacts. However, there are certain drawbacks of using the wet electrodes: the skin needs to be prepared by a skilled person; the gel dries over time, leading to signal degradation; and the removal of the gel electrodes is often painful or leads to skin irritation.

To overcome the limitations of wet electrodes, dry electrodes are being integrated into wearable devices such as watches and hand-held monitoring devices thus making them user-friendly. The mechanistic principle involved in the functioning of dry electrodes is important to understand and is different from that of wet electrodes, and is discussed in Section 2. Based upon this mechanistic principle, this chapter provides a brief overview of different types of dry electrodes that exist in the literature. Factors such as the electrode area, material, applied pressure, and skin hydration that affect the electrode performance are presented in Section 3. This is followed by a discussion of the existing methods for testing the performance of these electrodes in Section 4. Finally, the chapter summarizes the key findings about the factors that affect electrode performance along with suggestions for future directions to aid wearable electrode development.

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2. Electrodes for biopotential acquisition

A lot of attention has been paid to ECG because of the clinical significance of heart disease, the ability to diagnose arrhythmias from biopotential, and the utility of long-term monitoring of ECG for some conditions that only occur intermittently and may therefore be missed in the clinic. Clinically, ECG is captured using wet electrodes. Some wearable devices, such as the Holter Monitor, capture ECG using wet electrodes on the skin and provides an ambulatory electrocardiogram [2]. Such a wearable device’s benefits include a portable device that is worn for a short time (24–72 hours) to determine if the subject has any occasional cardiac arrhythmias, but they do face the challenges associated with wet electrodes mentioned earlier, such as being uncomfortable and leading to skin irritation. Because of this, there is a lot of research and development activity to implement an ECG system using dry electrodes. Although not an exhaustive list, two of these are described here. One such device is KardiaMobile by AliveCor, a US-based company [3]. It is a hand-held device and can detect atrial fibrillation (an irregular and rapid heart rate). It uses two 3 cm × 3 cm stainless steel electrodes and provides single-lead ECG. Dry electrodes are also being integrated into wearable devices such as watches. One such device is an Apple watch series 4. It also measures a single-lead ECG and consists of two electrodes. The button on the side of the watch (digital crown) made up of titanium serves as one of the electrodes, and the back crystal electrode on the back of the apple watch serves as another electrode. It consists of an ultrathin chromium silicon carbide nitride layer that is applied to the sapphire crystal. The wrist of the user is always in contact with the watch and the user needs to apply his fingertip across the crown, which creates a closed circuit, and this ECG can be recorded using the ECG app. Although dry electrodes are more user-friendly, they generally lead to poorer quality signals, due to the different mechanistic principles of the wet and dry electrodes, which are discussed in the following section.

2.1 Mechanistic principle of wet electrodes

2.1.1 Electrochemical reactions of electrode/electrolyte interface

The mechanistic principle of a wet electrode is based on an electrode-electrolyte interface’s electrochemical reactions. When a metal electrode comes in contact with an electrolyte solution, redox reactions may occur at the interface. The redox reactions occurring at the electrode/electrolyte interface are complex, but mechanistic models have been proposed that describe this interface as the double layer structure. The very first model was proposed by Helmholtz [4] and then modified by Gouy [5], Chapman [6], Stern [7], and Grahame [8], and followed by Bockris, Devanathan, and Muller with what is known as the BDM model [9]. Details for the double layer structure can be found elsewhere [10, 11] but are described here in brief in the context of the workings of a wet electrode by means of the double-layer structure at the interface of an Ag/AgCl electrode immersed in an electrolyte containing NaCl (Figure 2) [12].

Figure 2.

An example image of a typical wet electrode is shown. The zoomed-in view represents the electrode-electrolyte interface for a wet electrode depicting the double layer structure and an electric field of the interface, where IHP is the inner Helmholtz plane and OHP is the outer Helmholtz plane (Image modified from reference [12]).

The electrochemical reactions occur at the liquid electrode interface when the Ag/AgCl electrode is immersed in an electrolyte containing Na + and Cl- ions. First, the oxidation reaction occurs at the Ag electrode interface with the AgCl coating, which results in an excess of Ag + ions. The Ag + ions move into the AgCl and fill up the vacancies or move the adjacent Ag + ions to an interstitial site. Further, the Ag + ion interacts with the Cl- ions in the electrolyte. Hence it leads to the precipitation of AgCl at the interface of AgCl and the electrolyte. Also, there is a dissolution of AgCl, as the Ag + ions in AgCl prefer to be reduced. The Cl- ions get dissolved in the electrolyte, which leads to a higher concentration of Cl- ions at the interface than at the bulk of the electrolyte. Thus, to neutralize the negatively charged plane, the Ag electrode becomes positively charged, and these two planes form a double layer structure, and there appears a potential difference across the electrode/electrolyte interface. The electrochemical reactions can be described by the following eqs. [12]:

Ag→Ag++e−Oxidation
Ag++Cl−→AgClPrecipitation
AgCl→Ag++Cl−Dissolution
Ag++e−→AgReduction

The potential difference across the electrode/electrolyte interface causes the rearrangement of Cl– ions and leads to the orientation of water molecules. The water molecules align due to their dipole nature. Some Cl– ions are specifically adsorbed on the AgCl surface. There also happens to be the electrostatic attraction of Cl– ions by the positive charges of the Ag/AgCl electrode. This process can be described as the formation of planes, the inner Helmholtz plane IHP and the outer Helmholtz plane (OHP). The plane crossing the center of the water molecules is called IHP, and the plane crossing through the center of the aligned Cl- ions is known as OHP. The two positively (Ag/AgCl) and negatively charged planes (OHP) act as two plates of a parallel plate capacitor, and thus can be described as a capacitor. It is also known as the Stern layer. Due to the thermal driving forces and electrostatic attraction, there is a diffusion of the other Cl- ions in the electrolyte. The concentration of Cl- ions decreases exponentially; it is known as the diffusion layer. The diffusion layer ends when there is no potential gradient, which is represented by the bulk electrolyte. The double-layer structure is formed by the combination of the stern layer and diffusion layer. It can be considered the two layers connected in the series and can be described by the following expression, where CH is the Helmholtz layer capacitance, and CGC is diffusion layer capacitance (Gouy- Chapman):

1Cdl=1CH+1CGCE1

2.1.2 Equivalent electrical circuit for electrode/electrolyte interface

Electrically, the electrode/electrolyte interface and the skin can be represented by a combination of resistor and capacitor networks. The double layer is formed at the interface of two different phases. If there are any electrochemical reactions at the interface, they take place in the double-layer structure, leading to the faradaic current. Faradaic current is the current generated by the redox reactions at the electrode. The double-layer structure is represented by a capacitor (Cd), while the faradaic current flowing through it due to electrochemical reactions can be expressed as leakage current of the capacitor and is represented by a resistor in parallel (Rd) [13]. It is also known as charge transfer resistance and can be represented as Rct. Thus, in wet electrodes, due to this charge transfer phenomenon, resistive coupling is dominant. In this work, as shown in Figure 3, Rct is used to represent the charge transfer resistance, Cdl as the double-layer capacitance, and the resistance offered by the electrolyte is represented by Rg.

Figure 3.

An analogous electrical equivalent circuit of the wet electrode. Electrode-electrolyte interface is shown, where Rct and Cdc represent the charge transfer resistance and double-layer capacitance respectively. Rg represents the resistance in the gel/electrolyte. The stratum corneum is shown as a parallel combination of resistor (Rsc) and capacitor (Csc), and deeper tissue layers as a resistor (Rd).

2.1.3 Equivalent electrical circuit for skin

The electrode-skin interface plays an important role in measuring biopotential signals. The anatomical structure of skin comprises different layers. It can be broadly divided into the epidermis, dermis, and subcutaneous tissues. The topmost layer of the epidermis is the stratum corneum (SC), which consists of dead cells and is also referred to as the horny layer [14]. The outermost layer has the highest electrical resistance as it consists of dead cells. The layer beneath the epidermis is known as the dermis and mainly consists of blood vessels and sweat glands. Sweat ducts penetrate the stratum corneum, and as the sweat emerges, it results in a low-resistance parallel pathway. Sweat is considered a weak electrolyte and thus the flow of sweat across the duct walls leads to an increase in the stratum corneum’s hydration. Hence, this causes variation in the conductance of the skin. The skin can be represented electrically by equivalent capacitor and resistor configurations.

Electrical conduction inside the body is ionic. The stratum corneum contains sweat ducts and hair follicles that contain an ionic liquid and traverse the SC, therefore allowing electrical conduction across the SC. Hence it can be represented as a resistor. Stratum corneum consists of dead cells, and from an electric standpoint, it can be considered an insulator between the electrode (one plate of the capacitor) and the other living conductive tissues underneath it (another plate of the capacitor). Hence, it can be represented by a capacitor. The skin can be modeled as a capacitor and resistor in parallel. Several investigators have studied human skin’s electrical properties in response to AC signals and found skin impedance is of the order of 100 Ω at high frequencies and 10 kΩ–1MΩ at low frequencies (below 1 kHz) [15]. Figure 4 shows an equivalent circuit model where Rsc and Csc represent the resistance of the stratum corneum, and Rd represents the resistance of the deeper tissue layers.

C=ε0εrAdE2
Zc=1ωCE3

Figure 4.

An analogous electrical equivalent circuit for the metal-based dry electrode. Contact is represented by a capacitor (Cc), along with the stratum corneum as a parallel combination of resistor (Rsc) and capacitor (Csc), and deeper tissue layers as a resistor (Rd). The air gaps are due to the surface roughness of the stratum corneum and stiff metal-based dry electrodes. Cc comprises both the air gaps and the thin native oxide that comes in direct contact with the skin.

In wet electrodes, the electrolyte/gel helps in facilitating the electrochemical reactions and hydrates the stratum corneum, thus providing conductive ions that create an easier ionic path between the electrode and the skin below the stratum corneum [16]. The gel further helps in lowering the skin-electrode impedance by reducing Rsc and increasing Csc. However, the gel dries over time, which leads to an increase in Rg, Rsc, and a decrease in Cd, and Csc. In addition to this, Huigen et al. revealed that the main origin of the noise in the surface electrodes is due to the electrolyte-skin interface and is highly dependent on the electrode gel [17]. Thus, to overcome the limitations and challenges of wet electrodes, efforts have been made in the field of dry electrodes.

2.2 Mechanistic principle of dry electrodes

In contrast to resistive coupling in wet electrodes, dry electrodes are free of gel, and thus in the absence of any sweat has no ionic fluid coupling from the skin to the electrode, and the mechanistic principle is fundamentally different [18] and referred to as capacitive coupling.

In the capacitive coupling associated with dry electrodes no actual charge crosses the skin/electrode interface, as the metal electrode being an inert metal is difficult to oxidize or dissolve [14]. A displacement current exists as a result of capacitance at the interface and the signal gets capacitively coupled from the body to the sensing electrodes [19]. The heart causes immediate changes in the electric potential within the tissue, which is sensed by the metal electrodes. The metal electrode acts as one plate of a capacitor, with the deeper tissue layers as the other plate of the capacitor. The thin metal native oxide, air gaps, and the dry outer layer of the skin (stratum corneum) together act as a dielectric (the space between plates of a capacitor) [18]. Thus Cc varies as per Eq. (2), where the capacitance is directly proportional to the relative permittivity and area, and inversely proportional to the distance between two plates. Further, the capacitive impedance holds an inverse relation to the capacitance given by Eq. (3). Based upon this mechanistic principle of dry electrodes, different types of electrodes are reviewed in Section 3.

Where ε0 is the relative permittivity of free space, εr is the relative permittivity of a material, A is the area of the plate, and d is the distance between the two plates, ω = 2πf, ωis in rad/s, and f is the frequency (Hz).

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3. Types of dry electrodes

Dry electrodes can be broadly classified into two major categories: contact and non-contact electrodes. As the name suggests, contact electrodes are in direct contact with the skin and can be made from a range of materials, including metal, textile, and polymer electrodes. Metal-based electrodes use metal plates such as stainless steel, silver, titanium, and gold [16, 20, 21, 22]. These provide ease of manufacturing, integration, and use, but due to the lack of an electrolyte and flat hard surfaces, it is difficult to achieve good contact with dry and hairy skin. This results in an increase in resistance of the stratum corneum (Rsc) and a decrease in stratum corneum capacitance (Csc) and contact capacitance (Cc). It leads to high skin-electrode contact impedance, which leads to poor signal quality. Efforts have been made to embed the metal-based electrodes on a foam structure for ease of wearing [20, 23].

Several investigators have worked on fabricating dry electrodes based on the use of textiles in place of stiff metals [24, 25]. Textiles-based electrodes are made using conductive thread, steel, yarn, or silver coated on nylon fabric. These electrodes can be embedded within clothing and used as smart garments to sense biopotential signals [26, 27]. Due to flexibility, there is an increase in the effective skin contact area compared to the metal-based electrode, which decreases the Rsc and increases the Csc and Cc, thus lowering the overall impedance. But still, the challenges of high skin-electrode contact impedance remain the same due to the absence of an electrolyte.

To enhance elasticity, electrodes have been fabricated on polymeric substrates like polyethylene terephthalate (PET) and coated with conductive polymers such as MWCNT/PDMS [28]. Another polymer material, such as PEDOT, poly (3,4- ethylenedioxythiophene), is coated on metal electrodes to enhance the electrochemical performance, as it would lead to an increase in Cc [29]. But often, fabrication is complex, and in the long run, polymers usually flake off. Another type of dry contact electrodes includes microtips/pin based. The surface of the electrodes coming into contact with the skin consists of pins/spikes. The pin’s height is of the order of 150 μm, which is sufficient to penetrate the 20 μm thick stratum corneum. Thus this leads to a lower value of Rsc and a higher value of Csc. Hence, the pins directly come into contact with the fluids underneath the stratum corneum, which acts as an electrolyte, and eventually, the spike electrodes behave like wet electrodes. Though the application of pin electrodes can overcome the high impedance of the stratum corneum [30], their use can lead to skin irritation and infections as they come into contact with fluids underneath the skin. There are other types of electrodes, which do not make direct contact with the body, which thus can be categorized as non-contact electrodes [31]. These electrodes are worn on top of clothing, and can acquire biopotential signals through clothing [32, 33, 34, 35]. The addition of clothing further increases the contact impedance, thus increasing the Rsc and decreasing the Cc and Csc.

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4. Significant parameters for dry electrodes

In this section, several factors affecting dry electrode performance such as the electrode area, electrode material, skin hydration, and pressure are discussed. In addition to the high skin-electrode impedance of the dry electrodes, the skin-electrode impedance is variable with the applied pressure. Major factors that should be considered in the development of new electrode designs and their contribution to the skin-electrode interface have been summarized in Table 1.

Factors affecting electrode performanceRscCscCcNet effect on impedance
Electrode MaterialNo changeNo changeChangesChanges
Increase in Electrode AreaDecreasesIncreasesIncreasesDecreases
Increase in PressureDecreasesIncreasesIncreasesDecreases
Increase in Skin HydrationDecreasesIncreasesIncreasesDecreases

Table 1.

Different factors that affect the electrode performance and the equivalent impedance fitting parameters that will be affected corresponding to these factors are shown. Rsc represents the resistance of the stratum corneum, respectively; Csc and Cc represent the capacitances of the stratum corneum and contact.

4.1 Electrode area

Researchers have tried to overcome the challenges of dry electrodes by increasing the area of the electrode [28, 36]. Electrode-skin impedance plays a major role in the biopotential signal quality. With an increase in the electrode area, the Rsc decreases (Resistance is inversely proportional to the area), and both Cc and Csc increase, which leads to a decrease in Zc. Thus, it leads to an overall decrease in skin-electrode impedance with an increase in the area of the electrode. This approach of increasing the area to achieve lower skin-electrode contact impedance is useful and effective, however, it has a limited application in devices such as wearables where small electrode size is necessary for devices to be practical as wearables. This is particularly true for watch type devices, which require very small (<1cm2) electrodes.

4.2 Electrode material

Another aspect to improve the quality of biopotential signal acquisition is using different electrode materials [20, 37, 38]. The classical approach followed by researchers for characterizing different electrode materials is entirely empirical, and consists of carrying out experiments with different chosen materials followed by qualitative or quantitative comparison with wet electrodes (gold standard) [20, 37, 38, 39]. After that, the most satisfactory material based on the performance is selected. Some researchers have used the equivalent circuit electrode-skin interface impedance models and fitted them in the electrochemical impedance spectroscopy experimental data to characterize the wet and dry electrodes [40]. Typically, equivalent-skin interface models have been used to demonstrate the role of the size of electrodes on skin-electrode contact impedance, which clearly depicts the decrease in skin-electrode impedance with an increase in the electrode area. However, the role of electrode material (typically metal-based) in the skin-electrode interface is not reported. In one of the recent studies [18], efforts were made to understand the rationale behind a “metal-based electrodes’” material performance. The electrode-skin impedance dependence on the electrode material was investigated by developing a skin-electrode interface model that includes aspects of prior models and incorporates a model of the electrode material. The findings of this work suggested that the relative permittivity of the electrode material native oxide plays a significant role and a higher value leads to an enhanced capacitive coupling and thus a lower skin-electrode impedance. As per Eq. (2), a higher relative permittivity εrleads to increase in Cc which further leads to lower Zc, thus lower skin-electrode impedance. However, the skin properties Rsc ad Csc are not affected by the change in the electrode material, but Cc can be improved significantly. Thus, investigating dielectric properties, such as thickness, and relative permittivity of the native oxide, can be one of the approaches when selecting material for biopotential electrodes.

4.3 Skin hydration

The impedance of the stratum corneum is greatly affected by changes in skin hydration. According to one study, the stratum corneum resistance (Rsc) decreases by 14 times, and capacitance (Csc) increases by 1.5 times under hydrated conditions as compared to dry skin [41]. The increased capacitance (Csc) is attributed to the fact that hydration significantly affects the dielectric properties of the stratum corneum. As per previously reported findings, the relative permittivity of dry stratum corneum is 10 [41] and that of hydrated SC is 49 [42], which suggests an increase in the Csc. Moreover, hydrated skin leads to higher Cc, thus enabling enhanced capacitive coupling. The increased contact capacitance can be attributed to the improved contact between the electrode and stratum corneum [18]. Thus hydration plays an important role in lowering the stratum corneum impedance and is the most significant factor in achieving a low skin-electrode impedance.

4.4 Pressure

For dry electrodes, researchers have made efforts to understand the role of applied pressure on electrode performance [25, 36, 43]. An increase in applied pressure results in a lower skin-electrode impedance and this can be attributed to the increase in the effective contact area due to applied pressure. As per Eq. (2), the increase in the effective electrode contact area leads to an increase in both Cc and Csc. In addition to this, there is a decrease in Rsc as resistance is inversely proportional to the area. Changes in the applied pressure affect the skin-electrode impedance significantly, which further impacts the signal quality. Therefore, applied pressure should be accounted for during the testing of wearable devices.

4.5 Summary of effects

The factors discussed in the above section and the impact on each parameter of the circuit model, along with the net effect on skin-electrode impedance are summarized in the Table below.

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5. Testing method of dry electrodes

This section reviews the existing methods used for developing and testing the newly designed dry electrodes. The existing methods has been broadly classified into in vivo and in vitro type of testing. In vitro testing of dry electrodes refers to the evaluation of the performance of wearable electrodes across the human subjects, and in vitro testing refers to evaluating the electrode performance on a platform (aritifical skin) that has similar properties to that of human skin.

5.1 In vivo testing

Historically, the development and testing of electrodes are empirically performed by conducting experiments in vivo on human volunteers. This method of testing has two limitations: First, the properties of the human skin change over time, which leads to changes in the skin impedance, thus making it challenging to perform reproducible measurements [24]. Secondly, the properties of human skin differ for each person. Hence, the intra and inter-subject variability complicate the understanding of the performance of the dry electrodes in the acquisition of biopotential signals. Therefore, the performance of electrodes cannot be investigated and the role of various electrode parameters including electrode material, and applicability of the electrode design to a wider range of populations, cannot be clearly understood.

5.2 In vitro testing

To overcome the limitations of in vivo testing, some efforts have been made to fabricate a synthetic model of skin and tissue. These so-called phantoms simulate the electrical properties of tissues. The phantoms can be used as controlled benchtop testing platforms and can be used to facilitate the development of the electrode designs as there is no explicit control on human variables.

Figure 5 shows the electrical properties of two layers of the skin, which comprise the topmost layer known as the stratum corneum, and all the inner layers of the skin other than the stratum corneum, which are considered dermis/deeper tissue layers. The electrical properties of the stratum corneum represented in Figure 5 were investigated in one of the initial research works by Yamamoto et al. [44], where the stratum corneum was removed from a forearm of a human subject with a cellulose adhesive tape. The keratin layers of the stratum corneum were removed stepwise by stripping the skin with the cellulose tape and the electrical properties of the dermis and stratum corneum were studied. The impedance of the stratum corneum accounts for a major portion of the whole impedance measured at the skin, and it dominates in the low-frequency range between 1 Hz and 10 kHz. The region of interest in electrophysiological signals is low-frequency regions, and the stratum corneum impedance plays the dominant role in the low-frequency region. In this section, the existing phantoms have been broadly classified into single-layered and two-layered phantoms.

Figure 5.

Electrical properties of two layers of skin, where ρk and εk are average resistivity and dielectric constants of stratum corneum; ρc and εc are average resistivity and dielectric constants of inner layers (dermis) (Image reproduced from reference [44]).

5.2.1 Single-layered phantom

Skin phantoms, also known as artificial/synthetic skin, are made using different materials to mimic various kinds of skin properties. Skin properties can be broadly categorized into surface, mechanical, acoustic, optical, electrical, and thermal properties. For mimicking electrical properties, gelatinous substances and elastomers are mostly used [45]. Several researchers have fabricated skin phantoms using gelatinous substances such as gelatin [24] and agar [15, 46]. The agar-based phantom’s electric and dielectric properties can be adjusted by adding NaCl and polyethylene powder, respectively. Using this approach, Ito et al. fabricated the conventional electromagnetic phantom as a whole layer and used it to mimic electrical properties for higher frequency ranges (300 MHz–2.5 GHz) [47]. Later on, Yamamoto et al. used the above approach and further tuned electrical properties by adding the carbon microcoil to increase the relative permittivity of the phantom and thus fabricated a phantom that simulated electrical properties ranging between 1 MHz up to 2.5 GHz. However, the hydrous phantoms could not be further tuned to mimic the electrical properties for frequencies below 1 MHz [48]. Thus, the conventional approach using agar, DI water, NaCl, and polyethylene simulate the electrical characteristics of a high-content water tissue such as muscle and brain, and is incapable of simulating the electrical properties of the skin.

In another such work, Kalra et al. fabricated a single-layer phantom using oil in gelatin for simulating dielectric properties in a low-frequency range (20 Hz–300 kHz). However, the results are four orders of magnitude away from the desired skin dielectric properties in the low-frequency range [49]. A similar approach using gelatin was followed for fabricating a phantom for electrophysiology in the recent work of Owda et al. A single-layer phantom was made, and efforts were made to tune the electrical properties with different gelatin and NaCl concentrations [50]. The contact impedance profile was compared for the developed gelatin-based phantom and ex vivo porcine skin over 20–1000 Hz [50]. However, the impedance of human skin is much higher than that of porcine skin. For biopotential signals, the stratum corneum impedance is of interest as it dominates the skin impedance in the low-frequency range (1 Hz to 10 kHz) [51]. Therefore, a single-layer phantom approach does not include the effect of the outer layer of skin. Hence, there is a need for a two-layered skin phantom.

5.2.2 Two-layered phantom

One of the two-layered artificial skin phantoms using elastomers was developed by Nachman et al. [52] to study the mechanical properties such as friction of dry and moist skin. The model consisted of two different layers, a hydrophilic silicone-based top layer, and a polyurethane-based dermis. But, the materials were selected to mimic the elastic modulus properties [52]. Another two-layered artificial skin phantom demonstrating both electrical and mechanical properties was demonstrated by Liu et al. [53]. The skin phantom consisted of a gelatin membrane mimicking the epidermis, SU-8 photoresist mimicking stratum corneum, and microholes mimicking sweat pores and simulating the electrical properties of skin in the frequency range of (20 Hz–1 MHz). However, this frequency range is still too high of a frequency range for characterizing the biopotential electrodes.

Since biopotential signals lie in the range of 1 Hz–1000 Hz, and hydration plays a significant role in the skin-electrode impedance, none of the phantoms described above are capable of simulating the range crucial for modeling the hydration state. To fill this gap, in one of our previous works [54], a two-layered phantom was simulated, where the two layers represent the deeper tissues and stratum corneum. The lower layer of the phantom mimicking deeper tissues was fabricated using a mixture of polyvinyl alcohol cryogel (PVA-c) prepared with 0.9% W/W saline solution. The upper layer representing the stratum corneum was simulated using a 100 μm-thick layer fabricated by spin-coating a mixture of polydimethylsiloxane (PDMS), 2.5% W/W carbon black (CB) for conductance, and 40% W/W barium titanate (BaTiO3) as a dielectric. The hydration of the stratum corneum was modeled in a controlled way by varying the porosity of the phantom’s upper layer. Steps for fabricating the two-layered phantom are shown in Figure 6 [54]. The fabricated phantom was found to simulate electrical properties in the range of 1 Hz–1000 Hz, and matched well with the physiological skin impedance of human subjects. Moreover, with the novel approach of change of porosity, it provides a capability to simulate the hydration in a controlled manner and the phantom can be tuned as per skin ranges among different individuals.

Figure 6.

Fabrication steps of a two-layered phantom. Clockwise from top-left: A mixture of PDMS, carbon black, barium titanate; followed by removal of air bubbles in the vacuum chamber; spin coat mixture at 1000 rpm and 30 s; cure in an oven at 80°C for 2 hours; laser-cut holes; Peel off the layer; cast PVA solution; freeze (12 h at −20°C); thaw (24 h at RT). (Reproduced from reference [54]).

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6. Conclusion and future outlook

This chapter reviews the factors that affect wearable electrode performance, which are important to understand for the development of these electrodes. The mechanistic principle along with an electrical equivalent circuit for both wet and dry electrodes is discussed that explains the differences in the transduction mechanism of wet and dry electrodes. Various types of dry electrodes have been briefly covered to highlight the current state of art. Efforts made by researchers in developing different types of dry electrodes in a detailed manner can be found in the literature in recent review papers [55, 56, 57, 58]. Several factors such as the electrode area, skin hydration, electrode material, and pressure that substantially impact the performance of the dry electrodes and their contribution to the individual parameters of the electrical model are discussed. This review suggests that higher relative permittivity of the electrode material’s native electrode, an increase in electrode area, the application of pressure, and hydrated skin can help in achieving lower skin-electrode impedance. Thus, these factors can be used for the development of wearable biopotential electrodes, to improve high-quality biopotential signals. In addition to the consideration of the factors for the development of the electrodes, testing the wearable electrodes in a controlled way is crucial for evaluating their performance. Thus several research efforts to fabricate the skin phantoms simulating the electrical properties of the skin across a broad frequency range of 1 Hz–300 GHz are reviewed. Based on this review, phantoms simulating the electrical properties of the skin across their corresponding frequency range are shown in Figure 7. In addition to this, the skin phantom that is capable of simulating the electrical properties of stratum corneum in the frequency range crucial for biopotential signals (1 Hz–1000 Hz) along with the controlled hydration status is discussed, which can be used to model the interaction of dry electrodes.

Figure 7.

Summary of the fabricated skin phantoms across the different frequency ranges. Numbers in square brackets show the reference numbers.

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Acknowledgments

This research was funded by National Institute on Health, grant number R01 NR018301.

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Conflict of interest

The authors declare no conflict of interest.

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Written By

Krittika Goyal and Steven W. Day

Submitted: 02 December 2022 Reviewed: 23 March 2023 Published: 18 April 2023

© 2023 The Author(s). Licensee IntechOpen. This chapter is distributed under the terms of the Creative Commons Attribution 3.0 License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

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