1 Introduction

Healthcare monitoring electronics has attracted great research interest due to their ability to provide real time monitoring to prevent people suffer from chronic diseases, such as cardiovascular diseases [1, 2]. In general, the heart-related diseases, such as arrhythmias and related deaths, can be prevented by the diagnosis of electrocardiogram (ECG) signals [3]. The ECG signals can reflect the electrical activity of heart muscles [4], which have been widely used to observe the activity of human heart and evaluate the cardiac and cardiovascular functions [5]. In a conventional ECG system, the ECG signals are recorded by attaching the silver/silver chloride (Ag/AgCl) electrodes on the skin to capture the bio-potential signals and then convert the signals to electrical signals with digital filtering system [6]. The Ag/AgCl electrodes are normally used together with a conductive hydrogel and an adhesive layer to reduce the impedance between the electrode and skin [7]. However, the hydrogel and adhesive layer can gather hair and cause skin inflammation, and the drying of hydrogel over time might lead to low quality of signals and is not suitable for long-term monitoring [8].

A promising ECG electrode should fill the requirements of low contact impedance and motion artifacts, with expected long-time stability, as well as good flexibility to withstand some extensional strain [6]. Over the years, a range of dry electrodes has been investigated to avoid the problems associated with the hydrogel in wearable ECG monitoring systems [5, 9, 10]. To maintain intimate electrode/skin contact, a common strategy is to develop a flexible skin-like conductive film by mixing stretchable polymers with conduct additives, such as conductive polymer [11], silver nanowires [12], carbon nanotubes [13], and carbon black powder [14]. These dry electrodes solve the disadvantages of Ag/AgCl electrodes with skin irritation and long-term wear, but the fabrication processes of such electrodes are complicated, not systematic, and unsuitable for mass-production. Another strategy to fabricate dry electrode is based on textiles. Conductive textile electrode can be fabricated by depositing conductive materials, including PEDOT:PSS [15], silver paste [16], and graphene [17], on fabrics by printing technique such as screen printing [18], or inkjet printing [17, 19]. The printing technique, which has been widely used in wearable and flexible electronics [20, 21], offers a promising versatility for ECG monitoring.

Graphene, a two-dimensional carbon nanomaterial with a hexagonal honeycomb lattice, has excellent conductivity, good flexibility, and less affected with temperature changes [22], which makes it attract extensive attentions in materials science and biomedicine, and has great potential in healthcare monitoring [23]. Graphene has been widely used in electron and quantum dots [24], MEMS and microfluidics [25, 26], epidermal electronics [27], supercapacitors [28, 29], batteries [30-35] and flexible electronics [36, 37]. Graphene dry electrode produced by chemical vapor deposition (CVD) method and transfer process has been used for ECG monitoring [38, 39]. Celik et al. [38] transferred CVD synthesized graphene on the surface of Ag/AgCl electrodes to prepare electrodes. The prepared graphene is deposited on rigid substrates, and the long-term monitoring comfort needs to be improved. Lou et al. [39] prepared graphene electrodes on the PET substrate, which has good toughness. However, The PET substrate is not conducive to skin breathing. Moreover, the CVD and transfer processes are complex, which are not suitable for large scale fabrication. Graphene flakes can be exfoliated from nature graphite in large scale through liquid phase [40, 41], which can be further dispersed in solvents to form graphene-based conductive inks and deposited on textile for wearable ECG monitoring by printing technique [42]. The graphene–textile electrodes are normally prepared by coating textile with graphene oxide, following with a reduction process to convert the graphene oxide to graphene [8, 17]. However, the effective reduction of graphene oxide is normally need harmful chemicals [8], which limits its practical application for ECG monitoring. Moreover, there is a certain degree of graphene shedding during the reduction process, which affects the electrode quality.

In this work, we describe the fabrication of a dry electrode for ECG monitoring by screen printing the water-based graphene ink on a commercial cotton textile. Graphene nanoplatelets were chosen as the conductive materials due to its good conductivity and flexibility. Carboxymethyl Cellulose (CMC) was chosen as the surfactant for dispersing graphene because of its bio-compatibility and good stabilizability in water. The graphene ink can be successfully deposited on the textile substrate by screen printing process, following with a low drying temperature to remove water and form conductive graphene pattern. The graphene electrodes printed on textile present good contact with skin surface and comparable performance against conventional Ag/AgCl pre-gelled electrodes in ECG measurements. Moreover, the graphene-based textile electrodes showed excellent performance in recording of ECG signals even after 1000 bending cycles. The proposed approach paves the way for the development of bio-friendly, economical, and scalable dry electrodes for ECG monitoring.

2 Material and methods

The fabrication processes of graphene-coated textile electrode for ECG monitoring are illustrated in Fig. 1a. The graphene ink was prepared from graphene and Carboxymethyl Cellulose (CMC, Sigma-Aldrich) with a ratio of 5:1 by weight. Firstly, 0.9 g of CMC was dissolved in 50 ml of deionized (DI) water and magnetically stirred on a magnetic stirrer for 30 min at 550 r/min. Then 4.5 g of graphene nanoplatelets (M15, XG Sciences) was added into the CMC aqueous solution, followed by putting into Vertical planetary ball mill (MITR-YXQM-1L) and stirring for 15 min at the rate of 350 r/min to obtain the final graphene ink. A commercial cotton fabric with yarn count 40 × 40, density 133 × 100 and weighting 130 g/m2 was chosen as the substrate.

Fig. 1
figure 1

a Preparation process of screen printed graphene electrodes. b Experimental setup for the screen printing. c Photo of the fabricated graphene-coated electrode

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Screen printing was performed using an aluminum alloy screen printer with polyester screen printing mesh (150 mesh count) and designed pattern, as shown in Fig. 1b. The graphene ink was transferred to the screen printing template with a spoon and scraped at a 45° using a neutral blade, following by annealing in an oven at 80 °C for 30 min. For patterns with more than one printing layer, the printed patterns were dried at 80 °C for 30 min before the next printing cycle. Figure 1c shows the electrode sensing effective area with 20 mm × 20 mm, and the strip size is 5 mm × 10 mm. To obtain a stable signal, copper tape and conductive silver paste is applied around the contact surface of the narrow strip. Then, a layer of medical tape is applied on the copper tape. The sheet resistance of printed graphene patterns was measured using a handheld four-probe tester (M3, Suzhou Jingge Electronic). Each sample was measured 6 times, the average value and standard deviation of each data point were calculated. All tests were performed at room temperature. Impedance measurements were carried out using a LCR meter (HTOKI-IM3536) with scan frequency from 4 to 1000 Hz. The bending test was carried out by a home-build bending test machine. X-ray photoelectron spectroscopy (XPS) was measured on a Kratos Axis Ultra DLD XPS spectrometer (Kratos Analytical Ltd, Manchester, UK).

2.1 Results and discussion

To identify the quality of graphene, we measured scan electron microscopy (SEM) and XPS data of the graphene nanoplatelets, as shown in Fig. 2. It can be seen that the size distribution of graphene nanoplatelets is range from ~ 1 to 8 µm, which had an average size of 3.6 ± 1.3 µm (Fig. 2a, b). Figure 2c shows the XPS spectrum of graphene sample, which has a strong C 1 s peak, a small O 1 s peak, and almost invisible N 1s and S 2p peaks. Using the elements analysis, the atom weight of C, O, N, and S are 96.63%, 2.68%, 0.30%, and 0.39%, respectively. Moreover, the AFM data of typical graphene nanoplatelets show the thickness of graphene nanoplatelets is ~ 10 nm, which is similar to the value (8–10 nm) in the data sheet supported by the supplier.

Fig. 2
figure 2

a Typical SEM image of the graphene flakes, and b the size distribution chart (data captured from 18 SEM images). c XPS spectra of the graphene sample and d typical C 1 s binding of graphene sample

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Figure 3a demonstrates a typical scanning electron microscope (SEM, TESCAN MIR3) images of untreated cotton fabric, in which the microstructure of the cotton fibers is clearly visible and the fiber surface is relatively smooth. After the deposition process, the morphology of cotton fabric surface is clearly covered with graphene flakes, as shown in Fig. 3c. It can be seen that graphene flakes are successfully attached onto the surface of cotton fibers after screen printing deposition. Figure 3b, d exhibits the cross-sectional SEM images of cotton fabric before and after graphene deposition. The untreated cotton fiber is looser, while the cotton fibric become relatively tight and flat after deposition. This can be attributed to the binding ability of the binder we used in the ink.

Fig. 3
figure 3

The surface SEM images of a textile and c graphene-coated textile electrode and the cross-sectional SEM images of b textile and d graphene-coated textile electrode

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Figure 4a shows the variation of the sheet resistance of the graphene coated cotton fabrics with the number of printing repetitions. It can be seen that the sheet resistance of coated fabric was ~ 200 Ω/sq with one printing passes. There was a sharp decrease in sheet resistance by ~ 65% for the second printing pass. The sheet resistance of the coated fabric with three and four printing passes shows a small change from 46.6 to 42.2 Ω/sq. This could possibly be explained by coverage ratio of graphene flakes over fabric with increasing printing passes and form a continuous and uniform conductive film after three printed passes. The intrinsic impedance of graphene textile electrode with different printed passes is shown in Fig. 4b. It can be seen that the intrinsic impedance is almost constant for all frequencies for the various printing passes of graphene-coated fabric electrodes. The electrode with four printing passes shows an intrinsic internal resistance around 30 Ω, indicating the excellent conductivity of graphene film.

Fig. 4
figure 4

a Variation of sheet resistance with increasing number of printing passes. b The intrinsic impedance of graphene fabric with increasing number of printing passes. Variation with frequency in the electrode–skin impedance of Ag/AgCl and graphene electrodes with different printing passes: c dry condition and d wet condition

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Figure 4c, d shows the contact impedance of the electrode–skin with the frequency range from 4 to 1000 Hz at dry and wet conditions. The graphene–fabric electrodes were placed on a left wrist that is 2 cm apart and have not been pretreated by skin. For comparison, pre-gelled Ag/AgCl electrodes were attached on the same position as textile electrodes. Figure 4c exhibits the electrode–skin impedance of the Ag/AgCl and graphene-coated textile electrodes under dry condition. The contact resistance values of the Ag/AgCl and the graphene electrodes (one layer) at 4 Hz are 183 KΩ, and 1.25 MΩ, respectively. Figure 4d shows the electrode–skin contact impedance of the Ag/AgCl and graphene electrodes under wet condition (~ 1 ml of a natural emulsion, purchased at the CHANDO flagship store, milky white, used for ordinary skin care hydration, water retention, was applied to the contact interface of the skin and electrode). It can be seen that the skin-contact resistance of graphene-based fabric electrodes with different layers shows little difference with using emulsion, which indicates good coupling of the electrodes to the skin. Thus, the experiments mentioned below are carried out on the graphene electrodes with printing one layer.

The schematic of ECG signal acquisition, processing, and transmission is shown in Fig. 5a. The ECG signal obtained by electrodes is amplified by a preamplifier (IN129, TI) and then passed through a filtering module, including a high-pass filter with cut-off frequency at 0.05 Hz, a 50 Hz notch filter and a low-pass filter with cut-off frequency at 100 Hz. The obtained ECG signal is further amplified by the post amplifier with adjustable gain and sampled at 333 Hz with a microprocessor (STM32F407, ST) to digitalize. Finally, the signal is transmitted to the terminal device via the Bluetooth module (CC2541, TI). Figure 5b illustrates the experimental setup of ECG signal acquisition.

Fig. 5
figure 5

a The schematic of ECG signal acquisition, processing and transmission circuit board. b Photograph of ECG signal acquisition setup

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Figure 6a shows the ECG signals acquired from the Ag/AgCl (black line) and the graphene-coated textile (red and blue lines) electrodes during sitting, standing, and rotating conditions. It can be seen that the amplitude of ECG signal obtained from the graphene-coated textile electrodes is similar to that obtained from the Ag/AgCl electrode at sitting and stand conditions, which indicates the good skin-contact performance of the graphene textile electrodes. In contrast, both the Ag/AgCl and dry graphene textile electrodes show obviously baseline drift under rotating condition, while the wet graphene-coated textile electrode exhibits a relatively stable, low noise signal. Figure 6b shows the morphology of P-QRS-T waves obtained from Ag/AgCl and graphene electrodes. The P wave, QRS complex wave and T wave which are the characteristic waves of the ECG signal can be clearly distinguished. The signals are almost coincident except for a slightly larger in amplitude from Ag/AgCl electrode. The Pearson correlation coefficient of Ag/AgCl and graphene electrode is as high as 99.47%. These results indicate that the graphene-coated textile electrodes, even in dry condition, have comparable performance with the Ag/AgCl electrode for ECG signal acquisition.

Fig. 6
figure 6

a ECG signals from Ag/AgCl and graphene-coated textile electrodes for 20 s under different states (sitting, standing, rotating). b The P-QRS-T waves of ECG signals obtained from Ag/AgCl and graphene-coated textile electrodes

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To investigate the mechanical flexibility of the graphene-coated textile electrodes, cyclic bending test was carried out using a stepper motor bending test machine, as shown in Fig. 7a. The change of the resistance of graphene textile electrode is less than 2% over 100 bending cycles, as shown in Fig. 7b. Even at an increase cycle number of 1000, only a ~ 4.5% increase of the resistance is observed, which indicates the good flexbility of graphene-coated texitle electrode. Due to the excellent bending tolerance of graphene-coated textile electrode, it can successfully acquire ECG signals even after 1000 times bending cycles, as shown in Fig. 8. The measured ECG signals from graphene-coated textile electrodes after bending show the characteristic waves of the ECG signal at sitting, standing, and rotating conditions. These results prove that the graphene-coated textile electrode exhibits good flexible tolerance, which is important for its application in wearable ECG signal monitoring.

Fig. 7
figure 7

a The photography of graphene-coated textile under bending state. b Resistance of graphene textile electrode varies with the number of bends

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Fig. 8
figure 8

ECG signals from graphene-based dry and wet electrodes after 1000 bending cycles within 20 s under different states (sitting, standing, rotating)

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To further investigate the long-term performance of graphene electrode and commercial gel electrode, we used both electrodes to monitor the ECG signal for 24 h, as shown in Fig. 9a. It can be seen that the ECG signal obtained from the pre-gelled Ag/AgCl electrode and the graphene electrode showed typical ECG waves at both sitting and rotating state. After wearing 24 h, the ECG signal obtained from the pre-gelled Ag/AgCl electrode showed a decreasing in amplitude at sitting state and a large baseline drift under rotating state. This is due to the poor contact between the Ag/AgCl electrode and the skin with gel drying after 24 h. However, the ECG signal obtained from the graphene electrode showed changeless amplitude at sitting state and small baseline drift under rotating state. Moreover, after 24 h of wearing, the skin attached with graphene electrode was not significantly abnormal, while the skin attached with the Ag/AgCl electrode appeared red rash, as shown in Fig. 9b. These results further prove the suitability of the graphene-coated textile electrode for long-time ECG monitoring.

Fig. 9
figure 9

a The long-term recording of ECG signals obtained from pre-gelled Ag/AgCl and graphene electrodes over a period of 24 h. b The photography of forearm after 24 h wearing of Ag/AgCl and graphene electrode. The red line area indicates the electrode patch area

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3 Conclusions

In conclusion, we fabricated a graphene-based dry electrode on textile for its applications in ECG monitoring through screen printing technique. The ECG signals acquired from the graphene-coated textile electrode exhibited good quality, which were comparable or even better in wet condition with commercial pre-gelled Ag/AgCl electrodes. Moreover, the graphene-coated textile electrode showed excellent bending tolerance. ECG recordings from the graphene electrode displayed stable characteristics even after a thousand bending cycles. These results have laid an important foundation for the further development of textile electrodes for ECG monitoring. Considering this electrode can be fabricated with low-cost and mass production, it would be a promising tool for widely application in wearable bio-potential monitoring systems for the prevention of cardiovascular diseases.