Recent advances in electrode development for biomedical applications

Abstract

Elaborate electrodes that enable adhesion to the skin surface and effectively collect vital signs are necessitated. In recent years, various electrode materials and novel structures have been developed, and they have garnered scientific attention due to their higher sensing performances compared with those of conventional electrode-based sensors. This paper provides an overview of recent advances in biomedical sensors, focusing on the development of novel electrodes. We comprehensively review the different types of electrode materials in the context of efficient biosignal detection, with respect to material composition for flexible and wearable electrodes and novel electrode structures. Finally, we discuss recent packaging technologies in biomedical applications using flexible and wearable electrodes.

Introduction

Owing to the rapid development of information technology (IT) based on computer and network technology, IT has been applied in many fields, including the field of medical devices in recent years. In particular, the field of small and personal portable medical devices has become the central point for incorporating IT in the medical field[1, 2]. For future electronic medical devices, it is essential to strengthen relationships with demanding institutions such as hospitals and medical services; furthermore, relevant and complementary cooperation with industries are expanding significantly because of smarter electronic medical devices and their association with software and content [1].

In addition, the trend in personal disease management has changed significantly. As the paradigm of medical treatment shifts from treatment centered to prediction and prevention oriented, the importance of personal diagnostic technology is increasing [3]. According to the World Health Organization, the number of global cancer patients in 2018 was 18.1 million, of which one among five men and one among six women were diagnosed with cancer, and one among eight men and one among 11 women were expected to die [4]. Although treatment methods for cancer are constantly being developed, the mortality rate of cancer patients is still high owing to difficulties in the early detection of cancer and accurate diagnoses. To solve such difficulties, a next-generation convergence diagnosis system that can accurately diagnose various diseases is being developed through the combination of various technologies such as biotechnology, information communication technology, and conventional diagnostic technologies [5].

Hence, to materialize a personalized and prediction-oriented disease diagnosis, the fabrication of high-performance biosignal acquisition devices is essential. Biosignal acquisition devices are categorized into electrocardiogram (ECG), electroencephalogram (EEG), electromyogram (EMG), electrooculogram (EOG), skin conductivity (galvanic skin response), skin temperature, and photoplethysmography. Representative biosignals are shown in Table 1 [6].

Table 1.

Typical amplitude and frequency ranges of electrical biosignals

Reprinted from Ref. [6] with permission

Electrical biosignals	Amplitude (mV)	Frequency range (Hz)
Electromyogram (EMG)	0.5–5	2–500
Electrooculogram (EOG)	1–10	0.05–100
Electrocardiogram (ECG)	5 × 10− 4–5	Up to 100
Electroencephalogram (EEG)	2 × 10− 3–0.1	0.5–100
Evoked-potential/Even-related potential (EP/RP)	1 × 10− 4–0.2	1–300

The core technologies of biosignal acquisition devices are (1) sensor (detector and amplifier) technology to detect biosignals from a living body, (2) interfacial signal processing analysis technology, (3) printer and display technology to output the analyzed result, and (4) circuit technology considering safety issues. Among the core technologies, interfacial signal processing is one of the most crucial for obtaining minute biosignals. Most biosignal losses during its acquisition are from the interface condition between the electrode and the surface of the skin. Hence, soft and flexible electrodes that can offer intimate adhesion to the skin surface and effectively collect biosignals should be developed.

Herein, we briefly overview the types of soft and flexible electrode materials used in an efficient biosignal acquisition system. In Sect. 2, various electrode materials, including homogeneous and composite materials, are presented. The effective and novel structures of the electrodes are discussed in Sect. 3. In Sect. 4, packaging technologies for the construction of low-voltage circuitry are reviewed. In Sect. 5, future opportunities for commercializing high-performance soft and flexible electrodes in biosignal acquisition systems are discussed.

Materials

Homogeneous materials for soft and flexible electrodes

In clinical practice, a conductive paste or gel (electrolyte) is typically used to attach an electrode to the human skin. Conformal contact between electrodes and the human skin lowers the impedance of the skin, resulting in the efficient acquisition of biosignals. Electrodes used in such a manner are known as “wet electrodes.” In a wet electrode, a chemical reaction occurs where charges are transferred at the interface between an electrolyte without free electrons and an electrode without ions. Owing to such a chemical reaction, the neutral state of the interface between the electrode and electrolyte is compromised, and a potential difference from that of the surrounding area appears. This potential difference is known as the “half-cell potential.” The lower the half-cell potential, the better is the performance because the current passing through the interface between the electrode and electrolyte consumes less energy. Currently, silver/silver chloride (Ag/AgCl) electrodes are used widely in clinical practice owing to their ease of manufacture as well as safety for the human body owing to their low half-electrode potential. However, regardless of the material, the wet electrode presents several disadvantages owing to the attachment method. First, to attach the wet electrode on the human skin, dead skin cells at the desired location (skin abrasion) must be peeled off. In addition, gel application, impedance tuning on a device level and cleaning after the biosignal measurement are required. In addition, as the electrolyte dries over time, the impedance of the skin will increase, which is disadvantageous. Hence, it is not suitable for unprofessional persons to use wet electrodes in emergency cases.

Meanwhile, a dry electrode refers to an electrode that does not contain an electrolyte. The development of dry electrodes for the acquisition of biosignals have focused primarily on dry electrode materials, e.g., metals such as stainless steel, Ag/AgCl, and Ag. These electrodes can be utilized for the emergency monitoring of critically ill patients, brain–computer interfaces, etc. because they can be attached to the human skin surface quickly and easily. The utilization of dry electrodes has been limited to EMG, ECG, and EOG, where biosignals are significant. The acquisition of EEG signals using dry electrodes has not been extensively investigated because of their small signal amplitude. Herein, we provide an overview of dry electrodes used in biosignal acquisition.

Wei et al. reported a metal dry bioelectrode using red Cu (99.9 %), as shown in Fig. 1a [7]. The red Cu metal dry bioelectrodes were microstructured using laser micromachining technology. The synergistic effect of red Cu and the microstructure provided an enhanced biosignal acquisition of ECG from patients. Baek et al. reported the fabrication of Ti dry electrodes on a polydimethylsiloxane (PDMS) substrate for the measurement of ECG signals [8]. Owing to the utilization of PDMS, flexibility and biocompatibility were achieved; the Ti/PDMS dry electrodes exhibited long-term reliability. Meng et al. described the fabrication of a microdome array of Ni-based dry electrodes, as shown in Fig. 1b [9]. The performance of the Ni-based dry electrodes were comparable to that of a standard wet electrode (Ag/AgCl). The Ni-based dry electrodes were encapsulated with PDMS to enhance their biocompatibility, and the structural benefit from the microdome design resulted in the acquisition of high fidelity biosignals from patients. Jung et al. reported the fabrication of a flexible dry electrode based on carbon nanotubes (CNTs), as shown in Fig. 1c [10]. Commercially available CNTs were mixed with PDMS (concentration range: 1–4.5 %). It is noteworthy that the concentration and diameter of the CNT electrodes are important for the acquisition of ECG signals. In particular, CNT dry electrodes exhibited long-term stability and high robustness during the monitoring of motion and sweat.

Fig. 1.

a Schematic diagram of structure design of metal dry bioelectrode. Reproduced from Ref. [7] with permission. b Illustration of dry microdome electrode array. Reproduced from Ref. [9] with permission. c Schematic diagram showing fabrication process of CNT/PDMS composite electrodes. Reproduced from Ref. [10] with permission. d Chemical structure of PEDOT/PSS

The use of an electrode coated with a conductive polymer offers several advantages. For example, it is easy to increase the electrode surface area, control the hydrophilicity/hydrophobicity of the electrodes, and impregnate biomolecules. The Kipke group at the University of Michigan, Ann Arbor, demonstrated that when poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS, chemical structure is shown in Fig. 1d) doped with a surfactant (poly(oxyethylene) 10-oleyl ether) was electrochemically coated on an Au electrode surface, the electrode indicated an impedance that was 24 times lower than that of an unmodified electrode. In an in vivo insertion experiment, the impedance of the electrode modified with PEDOT increased by two to three times after 7 days of insertion; however, the impedance of the Au electrode increased by a factor of 0.5, indicating that the PEDOT-modified electrode had a significantly lower impedance and better biocompatibility than the Au electrode [11].

Composite dry electrode for flexible and wearable biomedical applications

Recently, researchers have employed printed/flexible/wearable dry electrodes and, more specifically, composite dry electrodes, for monitoring biomedical signals, e.g., ECG signals, which does not require the application of wet gel on the skin while providing conformal contact at the electrode–skin interface, better impedance with skin, less noise while in motion (e.g., sports) sweat secretions, etc. [12, 13]. Conventional Ag/AgCl wet gel electrodes cause inconvenience, instability, and infection to the skin and may be able to satisfy signal acquisition for wearable applications. In addition to their exceptional electrical properties, CNTs offer excellent mechanical properties; therefore, they are widely investigated as a promising candidate for a new generation of composite resources. However, the greater the dispersion and the weaker the interfacial bonding edge, the better is the efficacy of CNTs for supporting polymeric media [12]. Several techniques have been employed to improve adhesion at the interface. By adding 1 part per hundred resin (phr) of multiwall CNTs (MWCNTs) with styrene-butadiene copolymer, an increase in the modulus by 45 % and an increase in the tensile length by 70 % can be achieved. It was discovered that the conductivity increased by five orders of magnitude for 2 and 4 phr, forming a percolating network. Therefore, CNT-based elastomeric composites are suitable as a carbon-type supporting pitch for rubber constituents [12]. The ECG signal intensity of a dry electrode using MWCNT/PDMS, amalgamated as a conductive polymer, was discovered to be better than a marketable electrode with a wet Ag/AgCl layer [13, 14]. Sanchez et al. fabricated a CNT/polysulfone composite, which was reported to demonstrate adequate potential in screen-printed disposable electrodes to exhibit high electrochemical activity and mechanical stability in addition to allowing the incorporation of the HRP enzyme [15]. According to the operating principle, three types of dry electrodes were used: surface, invasive microneedle, and capacitive electrodes. Among them, capacitive sensors, which do not require any physical or conductive connection with the skin surface, indicate a higher potential for the mainstream market; however, data acquisition and fabrication techniques must be improved [16]. Conduction in the CNT/poly epoxy composite is primarily due to tunneling, and the insulating polymer is vital to the transport mechanism [17]. The CNT content and interfacing orientation are crucial to the temperature and conduction dependencies. (Fig. 2)

Fig. 2.

a Experiment setup for skin impedance measurement. b Setup for ECG measurements using MWCNT/PDMS composite dry electrode. c Comparison of dry and wet ECG signals obtained from wet Ag/AgCl electrode.

Reprinted from Ref. [13] with permission

Jung et al. fabricated a CNT–PDMS-based dry ECG electrode that can be easily applied to regularly used ECG devices to achieve optimum performance with different CNT orientations, thicknesses, and diameters [10]. The Gause factor of a CNT/PDMS composite with 18 wt% is 12.4, indicating its excellent pressure sensitivity for biometric sensors [18]. The advancement of novel stretchable conductors increases their potential to be applied in neuroscience as well as affects the development of new-generation wearable and implantable biosensing platforms with solution-processable and transportable energy substitutions [19]. Khan et al. established an MWCNT/PDMS composite layer on a thick layer of polyethylene terephthalate (PET) flexible substrate for bendable applications. Resistance change was observed in the sensors at convex and concave angles [20]. Lee et al. demonstrated a stretchable, elastomeric MWCNT/PDMS composite fabricated via mixing and an electrical percolation threshold method, which involved a significant decline in resistance was observed at 0.6 wt% of MWCNT fillers [21]. Using microcontact printing and casting mold techniques, PDMS was implanted into a conductive elastomer, whereas MWCNTs were extensively mixed into PDMS as conductive fillers using toluene as the standard solvent. The fabrication process was easy and inexpensive [22].

Two key methods, namely chemical and physical methods, are used to functionalize the fundamentally sluggish behavior exhibited by the CNT surface. A few chemical methods that can deliver covalent functional groups on top of the CNT surface include fluorination, carbene and nitrene addition, chlorination, hydrogenation, bromination, cycloaddition, and silanization.

Polymer coating around CNTs, use of different ionic surfactants, and the endohedral method are physical approaches that enable CNTs to be dispersed and improve the interfacial interactions between CNTs and the interface of the polymer. In chemical functionalization techniques, acids and chemicals are dispersed, which may damage the CNT interface [12].

In conclusion, composite dry electrodes have ample potential for achieving good conductivity, self-adhesiveness, mechanical flexibility and stretchability, biocompatibility, low interfacial impedance of the skin electrode, and excellent skin compliance. However, before CNT/PDMS dry electrodes can be applied extensively and commercialized, a few fundamental issues must be addressed: (i) their lower solubility and dispersion when combined with polymer resins; (ii) deprived CNT–polymer interfacial grip [23]. Using different types of mechanical approaches, including ultrasonics, stirring, shear mixing, ball milling, calendaring, and extrusion, a uniform dispersion may be achieved. Applicable methods and their fabrication processing conditions should be optimized to minimize mechanical damage to CNTs.

Electrode structure engineering

Electrode design is an important factor for improving sensing characteristics. An improved readout current response can be achieved using a desirable electrode structure. Instead of conventional flat-electrode deposition, periodic micro- or nanopatterned electrode structures have been proposed. Herein, we present a compilation of recent studies pertaining to novel electrode structures, including (1) microneedles, (2) porous surfaces, (3) nanomeshes, and (3) atomic phase structure engineering for high-performance biosensing applications.

A microneedle-structured electrode penetrates the skin, forms a contact interface between the microneedle and skin tissue, and removes the high impedance of the stratum corneum. The dry electrode of a microneedle reduces the motion artifact with a lower noise signal and accurately records the vital signs (Fig. 3a) [24, 25]. Needle-shaped electrodes have been fabricated using various processing techniques, such as photolithography with overcut photoresist [26–28], laser drilling, molding [29–31], and three-dimensional printing [32–35]. Using microneedle electrodes, the detection of biomedical signals from EMG, ECG, and EEG were achieved [36]. Furthermore, a real-time β-lactam sensing in vivo was conducted using a microneedle-based electrode (Fig. 3b) [37] In order to record an efficient biosignal, the form of invasive microneedle are fabricated. However, the absence of reliable fabrication method for a centimeter-scale microneedle array and an efficient protocol that describe the attachment of the microneedle array to the human skin surface must be investigated together.

Fig. 3.

a Equivalent schematics of flexible dry electrode (FDE) and flexible microneedle array electrode (FMAE). Reprinted from Ref. [24] with permission. b Post-use dose-response curve for microneedle-based biosensor. Reprinted from Ref. [37] with permission. c Illustration of fabrication process of graphene nanomesh structure. Reprinted from Ref. [44] with permission. d Measured electrochemical characteristics curves of relative changes in current peak at fixed potential vs. concentration of DNA detected. Reprinted from Ref. [45] with permission

To enhance the electrochemical sensing properties or to enlarge the active surface area, a novel porous structure has been presented [38, 39]. The porous structure of an Fc-CS/SWNTs/GOD film provided highly sensitive electrocatalytic properties for glucose detection limits with excellent biocompatibility [40]. In addition, a porous Na₂CO₃/PDMS structure that resulted in a body moving sensor operation with self-powered energy harvesting behavior has been proposed [41]. Recently, a nanomesh structure platform using graphene or transition metal dichalcogenides (TMDs) has been proposed [42, 43]. Performing photolithography using a block copolymer-based photoresist resulted in a nanomesh structure with periodic holes of a few nanometers. Owing to accurately controlled nanometer-sized holes in graphene or TMDs, significantly enhanced sensing performances, such as a high biosensing capability with a high surface-to-volume ratio, were achieved (Fig. 3c,d) [44, 45]. The most significant advantage of the nanomesh structure is its easy-to-fabricate pattern. As mentioned above, the self-assembled block copolymer offers elaborate hole patterns without additional or complex lithography processes. Although high surface-to-volume ratios have been achieved through nanomesh or porous structures, it is still necessary to develop reproducible and scalable manufacturing methods to implement practical applications.

Instead of electrode structure modification, the atomic phase structure engineering of the material was attempted. Through a phase transition from 2 H (trigonal prismatic coordination) to 1T (octahedral symmetry) of TMDs, excellent capacitive behavior, electrocatalytic performance for the hydrogen evolution reaction, and biosensing with high sensitivity have been achieved [46]. A 1T-phase WS₂ exhibited a highly sensitive and selective H₂O₂ biosensor with a wide range and low detection limits. Furthermore, the detection of three types of hormones (i.e., T3, T4, and PTH) was successfully demonstrated using 1T-phase MoS₂ on a polyimide-based flexible substrate [47]. Biosensing through atomic phase structure engineering has great potential, but stable phase structure transformation and stability thereof are still in the development level. (Fig. 4)

Fig. 4.

Types of wearable device package technologies and signal interconnections. a Fully soft package. Reproduced from Ref. [51] with permission. b Rigid components integrated on soft substrate. Reproduced from Ref. [56] with permission. c Rigid substrate covered with soft material. Reproduced from Ref. [57] with permission. d Interconnection implemented using serpentine metal structures. Reproduced from Ref. [54] with permission. e Kirigami structure. Reproduced from Ref. [62] with permission

Packaging

Packaging wearable sensor devices using mechanically flexible materials is crucial for achieving a long device lifetime, reliable health signal acquisition, and comfortable patient experience. The outlying package material should be able to support the integration of electronic elements. To achieve a comfortable, noninvasive wearable device for healthcare monitoring, the device should impose minimal restraint on the natural motion of the body. An application-specific design approach should be considered, depending on the functional requirements.

Strategies to integrate flexible, stretchable, and soft devices with electronic components can be primarily categorized into three types: fully soft, rigid components integrated on a soft substrate, and rigid substrates covered with a soft material. A fully soft package is inspired by the epidermal patch-type device, which comprises stretchable substrates with an elastic modulus similar to that of skin and exhibit conformal contact [48–51]. In the first soft wearable device, silicone-based materials (PDMS, Ecoflex, Solaris, etc.) were used as the substrate, and the electrodes for biomedical measurements were integrated into the wearable device [48]. With the development of technology, other materials that can be used as a supporting material or passivation layer such as poly(vinyl alcohol) films [52], PET [53], parylene [49], and polyimide [54] have been investigated. However, a fully soft device is limited in terms of circuitry integration, signal processing, power delivery, and wireless communication. The second approach, i.e., integrating rigid components on a soft substrate material, can support simple integrated circuits (ICs) with a certain degree of softness and flexibility [52, 55, 56]. This method involves the integration of commercial components or custom-designed ICs on stretchable metal interconnections of soft, flexible substrates. Although this approach can only support simple ICs and basic routing techniques, it provides more functionalities than a fully soft device. As the third strategy, a highly integrated printed circuit board is employed, where complex multifunctional ICs are packaged together with the electrodes using a polymeric material [57, 58]. Even though the rigid platform does not provide as much comfort as the epidermal patch, it can support advanced circuit functions such as high resolution, multifunctional signal processing, and high-speed wireless communication with reliability.

Additionally, metals that form a skin–electronic interface and signal interconnection should be considered. Gold and platinum are widely used as electrode contact materials owing to their biocompatibility and chemical inertness [48]. Copper is frequently used as a signal interconnect and antenna material owing to its low cost and low electrical resistivity characteristics [59]. To support flexible and soft substrates, these metal materials are designed as meandered, serpentine interconnects [60] as well as kirigami structures [61, 62] to endure mechanical stress.

Conclusions

In summary, we presented a comprehensive review of recent research progress pertaining to electrode development, classified into three categories: (1) homogeneous and composite materials, (2) structural aspects, and (3) packaging technologies.

First, (1) regarding the electrode materials aspect, the development of electrodes composed of homogeneous materials has been sufficiently progressed, in which has categorized into dry and wet electrodes. However, there are pros and cons for each of the dry and wet electrodes. In the case of the wet type, the electrode can be well-attached to human skin by using a conductive paste or gel, but it shows limitations in terms of biomedical aspects due to inconvenience of wetting on the body. On the other hand, the dry type has little physical discomfort, but it is difficult to accurately record biosignals owing to high impedance coming from the electrode, so technologies that complement those two issue are required.

Second, (2) regarding the electrode structure aspect, many progresses are intensively being conducted with the advancement of the electrode structure as exampled: microneedles, porous surfaces, nanomeshes, and atomic phase structure. However, it is still necessary to improve stability and reproducibility in the presented structures. Along with the verification of a new structure, tests for reproducibility and robustness should be accompanied. In addition, combining both comfort and adhesion that can be attached to the skin with higher compatibility is desirable, which enables effectively transmit human signals without noise.

Third, (3) regarding the packaging aspect, recent advances in the system packaging technology have demonstrated that soft electronic devices can be a perfect candidate for various wearable biomedical applications. Systems implemented on fully flexible soft material can provide patients with comfort, but hard circuit boards serve as a promising solution for high performance wearable systems. Hence, the researcher should thoroughly characterize the properties of the system prior to implementation and choose the solution that best fits the application. Packaging techniques are still developing with newly engineered materials and thorough clinical verification is necessary for practical realization. Interdisciplinary cooperation among researchers in material science, clinical medicine and engineering is important for commercialization and large-scale clinical studies are necessary.

To attain higher sensing performances in a practical system level compared with those of conventional electrode-based sensors, comprehensive effort should be made; co-development approach of material, structure, and packaging technology aspects lead to conformal contact between electrodes and the human skin such that the impedance of the skin can be reduced as well as offer new opportunities for realizing next-generation biomedical devices.

Declarations

Conflict of interest

The authors declare that they have no conflict of interest.

Ethical approval

This article does not contain any studies with human participants or animals performed by the author.