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Experimental investigation on surface wettability of copper-based dry bioelectrodes
Sensors and Actuators A 244 (2016) 237–242
Contents lists available at ScienceDirect
Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna
Experimental investigation on surface wettability of copper-based dry bioelectrodes Wei Zhou ∗ , Wei Liu, Shaoyu Liu, Guobiao Zhang, Zhijia Shen Department of Mechanical & Electrical Engineering, Xiamen University, Xiamen 361005, China
a r t i c l e
i n f o
Article history: Received 27 January 2016 Received in revised form 10 April 2016 Accepted 19 April 2016 Available online 22 April 2016 Keywords: Dry bioelectrode Superhydrophobicity Laser micromilling Surface microstructure
a b s t r a c t Laser micromilling technology was used to fabricate a dry bioelectrode with a surface microstructure array, which was then treated with stearic acid ethanol solution. A superhydrophobic dry bioelectrode was obtained. Based on the scanning electron microscopy results, the formation of the dry bioelectrode with a surface microstructure array had been discussed. Later, the contact angle on the bioelectrode surface was analyzed by changing the parameters such as laser processing parameters, spacing of microstructure, and the working media. The results show that the contact angle of the bioelectrode surface treated with stearic acid ethanol solution was approximately 151◦ , and this value was compared with the result for bioelectrodes that were not treated or placed in air. The laser power and number of scans had significant influences on the contact angle, while the scanning speed had only a slight influence. When the spacing of microstructure was in the range of 0.1–0.5 mm, the contact angle increased with a decrease in the spacing of microstructure. In addition, the contact angle on the bioelectrode with deionized water or saline water as the working media was greater than 150◦ . Our developed dry bioelectrode presented good hydrophobic properties to help to realize a stable measurement system over a long time, which has many applications in the detection and measurement of biological signals. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Biomedical electrodes can convert the ionic potential generated by electrochemical activities into an electronic potential of a measurement system, and this system is widely used in modern clinical detection and biomedical measurements including electrocardiographs (ECG),electroencephalograms(EEG), electromyographs (EMG), electrical impedance images (EIT), and measurements of gastric electrical activity and nerve potential [1–3]. Biomedical electrodes are usually composed of metal materials, silicon, or polymer as a base material, which is covered by a conductive metal layer to obtain good electrical conductivity. In the measurement process, biomedical electrodes are in direct contact with human skin to inject driving currents and receive voltage measurements to achieve information exchange and transmission. Thus, the contact resistance of skin electrodes must be reduced to obtain stable physiological signals. Biomedical electrodes must also be nonpoisonous and harmless and avoid adverse physiological reactions such as allergies [4,5]. In addition, a flexible substrate and a foam
∗ Corresponding author at: Department of Mechanical & Electrical Engineering, Xiamen University, Xiamen, 361005, China. E-mail addresses: [email protected], [email protected] (W. Zhou). http://dx.doi.org/10.1016/j.sna.2016.04.044 0924-4247/© 2016 Elsevier B.V. All rights reserved.
backing structure in bioelectrodes are also used to reduce the problem of motion interference during the measurement process. Until now, the developed biomedical electrodes are mainly traditional silver/silver chloride electrodes, microneedle electrodes, textile flexible electrodes, flexible substrate electrodes, foam structure electrodes, and insulation dry electrodes [6–8]. In recent years, biomedical electrodes technology has attracted increasing attention from researchers. The design and manufacturing technology of biomedical electrodes has made rapid progress. Nowadays, biomedical electrodes can be divided into two types: dry and wet bioelectrodes. Compared with wet bioelectrodes, i.e., Ag/AgCl electrodes, dry bioelectrodes are attractive because there is no need for an electrolytic gel for the bioelectricity impedance technique, which can effectively overcome the limitations of wet electrode technology and show excellent signal collection performance [3,4]. In particular, a dry bioelectrode with surface microstructure arrays has been developed for its good electrodeskin contact interface to prevent high impedance. Thus, a dry bioelectrode can be used in the long-term recording process without conductive gel and skin preparation, which also has several obvious advantages such as lower impedance variations, smaller electrochemical noise, better stability, and convenient use. Several fabrication methods for bioelectrodes, micromachining such as wet etching processing [9–12], reactive ion etching [13,14], the casting
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Fig. 1. Structural design of dry bioelectrode with a novel construction: 1-Back side of electrode back; 2-Front side of electrode; 3-Shielding wire; 4-Conductive silver glue; 5-Metal electrode core with surface microstructures arrays; 6-Foam backing material.
method [15], and 3D printing [16] have been developed to measure the physiological signal. Up to now, much research work about the fabrication and performance of bioelectrodes have been reported. The wet-etching technique was first proposed to fabricate bioelectrodes with surface micro needle structures [9–12]. Later, Griss et al. [13,14] developed the deep reactive ion etching method to fabricate spike electrodes using silicon that was subsequently covered with a Ag/AgCl double layer to reduce electrode interface noise. Ng et al. [15] developed the vacuum casting method to fabricate microspike dry EEG electrodes, characterized the sensing performance in terms of the impedance level and stability, and showed a much higher efficiency in EEG measurements. Salvo et al. [16] develop the 3D printing technology to fabricate electrodes with 180 conical needles. The ECG and EEG were then measured compared with planar wet Ag/AgCl electrodes. Wang et al. [17] developed two combined methods to fabricate silicon-based dry electrodes for measuring biological signals. Lopez-Gordo et al. [18] review current approaches to develop dry EEG electrodes with nano-, micro, and millineedles for clinical and other applications, including information about measurement methods and evaluation reports. Recently, Liu et al. [19] proposed a micromolding technology to fabricate a three-dimensional microneedle electrode arrays. Electrical impedance and polarization voltage experimental results showed that the electrode chips exhibited great electric characteristics. In the present study, laser micromilling technology was used to fabricate a metal dry electrode with a surface microstructure array, which was then treated with stearic acid ethanol solution. The contact angle of the bioelectrode was experimentally investigated by the varying the laser processing parameters, the spacing of the microstructure, and working media.
Fig. 2. Optical image for laser micromilling setup of dry bioelectrode with surface microstructures arrays. Table 1 Specifications of characteristic parameters of the used fiber laser system. Characteristic
Parameter range Process conditions
Unit
Wavelength Nominal average output power Pulse duration Repetition rate Beam quality (M2 ) Incident beam diameter Focused diameter The number of scans Laser output power Scanning speed
1055–1070 19–21 90–120 20–200 <1.1 6–9 24.3–37.3 0–50 0–0 0–1000
nm W ns kHz
1064 20 100 20 1 7 31.5 5, 10, 20, 50 14, 16, 18, 20 250, 500, 750, 1000
mm m / W mm/s
The specifications of the characteristic parameters of the fiber laser system are given in Table 1. Moreover, a linear motor was used to move the laser head in the z-direction so that the laser beam could be focused on the surface of the electrode sample materials by setting the focal length to 163 mm. To fabricate the bioelectrode with surface microstructures arrays, the cross machining route and the loop multiple-pass reciprocating scanning strategy were predetermined in this study [20–22]. During the laser micromilling process, the surface material can be removed layer by layer through repeating the scan sequence. After the laser micromilling process, the dry bioelectrodes were cleaned with diluted hydrochloric acid and ethanol in an ultrasonic bath for 5 min to remove surface oxides and any organic substance. Then the dry bioelectrode was completely impregnated in 2 g/L stearic acid ethanol solution for 2 h at 40 ◦ C. Finally, the dry bioelectrode was clean with deionized water and dried in a baking box. The surface morphology of the dry bioelectrodes was observed through a scanning electron microscope (SEM) (Hitachi SU70, Japan) with energy-dispersive X-ray measurements (EDX). .
2. Experimental procedure 2.2. Contact angle test setup 2.1. Design and fabrication of dry bioelectrode Fig. 1 shows the schematic diagram of structural design of dry bioelectrode with a novel construction. The novel designed bioelectrode is composed of metal electrode core, foam backing material, conductive silver glue and shielding wire. Optical image for laser micromilling setup of dry bioelectrode is shown in Fig. 2. In this study, a prototype pulsed fiber laser (IPG, No: YLP-1-100-20-20CN, Germany) was used as the fabrication laser. The laser was set to produce 100-ns pulses with a 1064-nm central emission wavelength at a repetition rate of 20 kHz. The nominal output power of the laser was 20 W, which was sufficiently higher than the ablation threshold of the bioelectrode sample materials (red copper).
The contact angle () measurement was carried out with an automatic contact angle meter (Model: DCAT21, Dataphy Instrument Co., Ltd, Germany). The test system was composed of a syringe pump, a CCD camera, a light source, a sliding table, and a computer analyzing system, as shown in Fig. 3. In this study, the static contact angles were actually obtained by sessile drop measurements, where a drop was deposited on the surface, and the value was obtained by a goniometer [23]. To measure the contact angle, a syringe pump was first used to inject a controlled volume of ∼3 L droplet of the working media on the sample surface through the syringe. After the droplet remained on the sample surface for about 30 s, images of the droplet were taken using the CCD camera.
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surements were estimated to be 1◦ . All the measurements were performed at room temperature (approximately 25 ◦ C). 3. Results and discussion 3.1. Surface microstructure of the bioelectrode
Fig. 3. Optical image of contact angle test setup for bioelectrode.
Finally, three images were analyzed using the supplied software to determine the contact angle. For each sample, the contact angle measurement was repeated on five randomly selected areas on the surface of the bioelectrode [24]. The average value of their contact angles was taken as the static contact angle of the samples. The standard deviation and the average error for contact angle mea-
In previous study, different surface micro/nanostructures have been fabricated using laser micromachining process [25,26]. The laser micromachining process can be simply defined as a material removal process by heating the surface material with a laser beam, which includes four processing stages of surface material heating, surface material melting, material evaporation, and ionization as well as evaporation spray and fluid ejection [27,28]. When the fiber laser emitted a collimated beam of white light to continuously heat the material surface, surface thermal energy accumulated quickly, resulting in a significant increase in the temperature of the material surface. After the temperature of the material surface was increased to slightly lower than the melting point of the material, the liquid phase damaged the material surface. Subsequently, with continuous irradiation of the laser beam, metal steam began to partially ionize to form the low-density plasma, which was helpful to further absorb laser energy. Because of the pressure of the vaporizing expansion and plasma, the surface liquid material gradually evaporated to induce the ejection phenomenon to produce many grooves.
Fig. 4. Typical images of the microstructure array on the surface of the dry bioelectrode:(a) Optical image; (b) SEM image; (c) Cross-section view; (d) Top view.
Fig. 5. Contact angle on the bioelectrode surface under different conditions: (a) untreated, (b) after one month in air, and (3) treated with stearic acid ethanol solution.
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7
7 C O
Weight (%)
6
6
5
5
4
4
3
3
2
2
1
1
0
a
b
c
0
Condition Fig. 6. The proportion of carbon and oxygen by weight of the bioelectrodes: (a) untreated, (b) after one month in air, and (3) treated with stearic acid ethanol solution.
In particular, a recasting phenomenon readily occurred to produce the recast layer structure to deposit the unscanned area between the process spacings. Thus, a small quantity of surface material was removed from the laser-scanned area, and the recast layer started to form around the unscanned area. When the number of scans was gradually increased, more recast layer was continually deposited and solidified on the unscanned area. Ultimately, the recast layer was deposited and combined on the unscanned area to form a microstructure array with different geometrical dimensions on the surface of the sample material. Typical images of the microstructure array on the surface of the dry bioelectrode are shown in Fig. 4 The surface microstructure array of the dry bioelectrode was successfully fabricated using the recast layer in the laser micromilling process.
3.2. Contact angle on the bioelectrode surface under different conditions Fig. 5 shows the contact angle on the electrode surface under different conditions. The bioelectrode with 0.1 mm spacing of surface microstructure was fabricated with a scanning speed of 500 mm/s, a laser power of 20 W, and 20 scans. For the untreated bioelectrode, water drops spread out immediately on the bioelectrode surface, so the contact angle was 0◦ . When the bioelectrode was placed in the air for one month, the contact angle was 140◦ . This indicated that the bioelectrode showed hydrophobic properties. Several previous studies also shows that micro/nano-structured metal surface become superhydrophobic from initial hydrophilic in air over time [29–31]. In particular, the contact angle was 151◦ when the bioelectrode was treated with stearic acid ethanol solution. Thus, a bioelectrode with superhydrophobicity could be obtained. Fig. 6 shows that the proportion of carbon and oxygen by weight changed significantly. The contact angle increased with the carbon weight. However, the contact angle did not change with the oxygen weight, because the C C bond influences the surface polarity. A higher C C value indicates a more nonpolar surface, which tends to hydrophobicity [32,33]. Compared with a copper sheet treated with stearic acid ethanol solution ( = 103◦ ), the contact angle of the bioelectrode with a surface microstructure array fabricated using the laser micromilling technique increased greatly. These results indicated that the surface microstructure array could change the state of contact with water droplets. A bioelectrode with surface hydrophobic properties could be obtained.
Fig. 7. Contact angle of bioelectrodes fabricated with different laser process parameters.
3.3. Effect of laser processing parameters on the contact angle Fig. 7 shows the contact angle of a bioelectrode fabricated with different laser processing parameters. With a laser output powers of 14W–20W, the contact angle increased with the increasing laser output power. After the laser output power increased to above 18 W, the contact angle reached a stable value of around 150◦ . When the scanning times were increased, with values of 5, 10, 20, and 50 times, the contact angle increased, as shown in Fig. 7b. When the scanning speed was less than 750 mm/s, the contact angle changed little. However, a slightly decreasing trend was observed when the scanning speed increased to 1000 mm/s, as shown in Fig. 7c. These results can be attributed that the surface morphology and size (such as height and spacing) of the microstructure array on the bioelectrode surface changed with different laser processing parameters [21]. Therefore, reasonable laser processing parameters, including
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Fig. 8. SEM image of bioelectrodes with different microstructure spacings.
0.1 mm and 0.2 mm. When the microstructure spacings increased to 0.5 mm, the contact angle deceased to 146◦ . Therefore, we conclude that the microstructure spacing can affect the value of the contact angle. Furthermore, all contact angles of bioelectrodes were above 146◦ , indicating good hydrophobic properties. For the small microstructure spacing, the droplets were lifted from the electrode surface. The contact angle was thus increased. Thus, the developed bioelectrodes with smaller microstructure spacing lift the droplets to exhibit the Cassie state [34,35]. However, with increasing microstructure spacing, the droplets can easily enter the gaps in the microstructure. The contact status also gradually shifted from the Cassie state to the Wenzel state, leading to a smaller contact angle. 3.5. Different working media on the bioelectrode surface Fig. 9. Contact angle of bioelectrodes with different microstructure spacings.
the laser output power, scanning time, and scanning speed, should be selected to fabricate bioelectrodes with much larger contact angles. 3.4. Effect of microstructure spacing on the contact angle Figs. 8 and 9 show SEM images and contact angles of bioelectrodes with different microstructure spacings. The contact angle increased with a decrease in the microstructure spacing. The contact angle can be above 150◦ when the microstructure spacings are
Because dry bioelectrodes are in direct contact with the related working medium during long-term measurement processes, it is important to study the contact angle of the electrode surface with different working media [16]. In particular, sweat is easily produced on the skin because of the body’s metabolic activity, which seriously affects the measurement results. Therefore, deionized water, 3 g/L saline water (similar to the sweat of sodium chloride concentration), and glycerin were selected as working media for the contact angle measurement. The results show that the contact angle of deionized water and saline water on the electrode surface could reach above 150◦ , indicating superhydrophobic properties. The contact angle for glycerin was 142◦ . In general, the dry electrodes showed good hydrophobicity and oleophobic properties, as
Fig. 10. Deionized water, saline water and glycerin droplets on surface of bioelectrode: (a) Optical images; (b) Contact angle value.
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shown in Fig. 10. When the dry electrode is in contact with deionized water, saline water, and glycerin, a larger contact angle can be obtained to maintain good contact conditions between the electrode and skin. Thus, the stability of the measurement process with a dry bioelectrode can be improved. 4. Conclusions The laser micromilling technology was used to fabricate dry bioelectrodes with a surface microstructure array, which were then treated with stearic acid ethanol solution. The contact angle of a dry bioelectrode could reach 151◦ , indicating superhydrophobic properties. After discussing the formation of the dry bioelectrode with a surface microstructure array, the contact angle on the electrode surface was analyzed by changing the parameters, such as through laser processing parameters, spacing of the microstructure, and the working media. The laser output power and the number of scans exhibited a great influence on the contact angle, but the scanning speed did not have much influence. The dry bioelectrodes with microstructure spacings of 0.1 mm and 0.2 mm presented good hydrophobic properties. In addition, the contact angles on the bioelectrode with deionized water and saline water as the working media were greater than 150◦ . Our developed metal dry bioelectrode presented a good hydrophobic properties to help realize stable measurements over a long time and have wide application prospects in the detection and measurement of biological signals. Acknowledgments This work was supported by the National Natural Science Foundation of China (Project No. 51475397) and the Fundamental Research Funds for Central Universities, Xiamen University (Project No. 20720160079). In addition, the supports from Open Fund of Shanghai Key Laboratory of Digital Manufacture for Thin-Walled Structures (No. 2015004) are also acknowledged. References [1] Y.M. Chi, T.P. Jung, G. Cauwenberghs, Dry-contact and noncontact biopotential electrodes: methodological Review, IEEE Rev. Biomed. Eng. 3 (2010) 106–119. [2] E. Spinelli, M. Haberman, Insulating electrodes: a review on biopotential front ends for dielectric skin-electrode interfaces, Physiol. Meas. 31 (2010) s183–s198. [3] M. Bodenstein, M. David, K. Markstaller, Principles of electrical impedance tomography and its clinical application, Crit. Care Med. 37 (2009) 713–724. [4] A. Searle, L. Kirkup, A direct comparison of wet, dry and insulating bioelectric recording electrodes, Physiol. Meas. 21 (2000) 271–283. [5] J.Y. Baeka, J.H. An, J.M. Choi, Flexible polymeric dry electrodes for the long-term monitoring of ECG, Sens. Actuators A 143 (2008) 423–429. [6] A. Gruetzmann, S. Hansen, J. Muller, Novel dry electrodes for ECG monitoring, Physiol. Meas. 28 (2007) 1375–1390. [7] C.T. Lin, L.D. Liao, Y.H. Liu, Novel dry polymer foam electrodes for long-term EEG measurement, IEEE T. Bio-Med. Eng. 58 (2011) 1200–1207. [8] W. Zhou, W. Liu, Q.F. Qiu, Development, fabrication, and applications of biomedical electrodes (in Chinese), Chin. Sci. Bull. 60 (2015) 1352–1360. [9] P.K. Campbell, K.E. Jones, R.J. Huber, K.W. Horch, R.A. Normann, A silicon-based three-dimensional neural interface-manufacturing processes for an intracortical electrode array, IEEE Trans. Biomed. Eng. 38 (1991) 758–768. [10] Q. Bai, K.D. Wise, D.J. Anderson, A high-yield microassembly structure for three-dimensional microelectrode arrays, IEEE Trans. Biomed. Eng. 47 (2000) 281–289. [11] C.O. Mahonya, F. Pini, A. Blakea, C. Webstera, J.O. Briena, K.G. McCarthy, Microneedle-based electrodes with integrated through-silicon via for biopotential recording, Sens. Actuators A 186 (2012) 130–136. [12] N. Wilke, C. Hibert, J.O. Brien, A. Morrissey, Silicon microneedle electrode array with temperature monitoring for electroporation, Sens. Actuators A 123–124 (2005) 319–325. [13] P. Griss, P. Enoksson, H.K. Tolvanen-Laakso, P. Merilainen, S. Ollmar, G. Stemme, Micromachined electrodes for biopotential measurements, J. Microelectromech. S. 10 (2001) 10–16. [14] P. Griss, H.K. Tolvanen-Laakso, P. Merilainen, G. Stemme, Characterization of micromachined spiked biopotential electrodes, IEEE T. Bio-Med. Eng. 49 (2002) 597–604.
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Biographies Wei Zhou received his B.S. degree in mechanical engineering from Xiangtan University, Xiangtan, China, in 2005. In 2010, he received his Ph.D. degree in mechanical engineering from South China University of Technology, Guangzhou, China. From 2010–2012, he was a postdoc researcher at Sun Yat-sen University. Now he is an associate professor at Xiamen University. His current research interests focus on design, fabrication and performance evaluation of biomedical device. Wei liu starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2014. Her research interests are in the area of measure methods of different bioelectrodes. Shaoyu Liu starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University 2014. His research interests are in microfabrication technology of biomedical device. Guobiao Zhang starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2013. Her research interests are in the area of modeling and simulation of microfabrication process of bioelectrode. Zhijia Shen starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2013. His research interests are in novel measurement technology for ECG and EMG.
Contents lists available at ScienceDirect
Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna
Experimental investigation on surface wettability of copper-based dry bioelectrodes Wei Zhou ∗ , Wei Liu, Shaoyu Liu, Guobiao Zhang, Zhijia Shen Department of Mechanical & Electrical Engineering, Xiamen University, Xiamen 361005, China
a r t i c l e
i n f o
Article history: Received 27 January 2016 Received in revised form 10 April 2016 Accepted 19 April 2016 Available online 22 April 2016 Keywords: Dry bioelectrode Superhydrophobicity Laser micromilling Surface microstructure
a b s t r a c t Laser micromilling technology was used to fabricate a dry bioelectrode with a surface microstructure array, which was then treated with stearic acid ethanol solution. A superhydrophobic dry bioelectrode was obtained. Based on the scanning electron microscopy results, the formation of the dry bioelectrode with a surface microstructure array had been discussed. Later, the contact angle on the bioelectrode surface was analyzed by changing the parameters such as laser processing parameters, spacing of microstructure, and the working media. The results show that the contact angle of the bioelectrode surface treated with stearic acid ethanol solution was approximately 151◦ , and this value was compared with the result for bioelectrodes that were not treated or placed in air. The laser power and number of scans had significant influences on the contact angle, while the scanning speed had only a slight influence. When the spacing of microstructure was in the range of 0.1–0.5 mm, the contact angle increased with a decrease in the spacing of microstructure. In addition, the contact angle on the bioelectrode with deionized water or saline water as the working media was greater than 150◦ . Our developed dry bioelectrode presented good hydrophobic properties to help to realize a stable measurement system over a long time, which has many applications in the detection and measurement of biological signals. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Biomedical electrodes can convert the ionic potential generated by electrochemical activities into an electronic potential of a measurement system, and this system is widely used in modern clinical detection and biomedical measurements including electrocardiographs (ECG),electroencephalograms(EEG), electromyographs (EMG), electrical impedance images (EIT), and measurements of gastric electrical activity and nerve potential [1–3]. Biomedical electrodes are usually composed of metal materials, silicon, or polymer as a base material, which is covered by a conductive metal layer to obtain good electrical conductivity. In the measurement process, biomedical electrodes are in direct contact with human skin to inject driving currents and receive voltage measurements to achieve information exchange and transmission. Thus, the contact resistance of skin electrodes must be reduced to obtain stable physiological signals. Biomedical electrodes must also be nonpoisonous and harmless and avoid adverse physiological reactions such as allergies [4,5]. In addition, a flexible substrate and a foam
∗ Corresponding author at: Department of Mechanical & Electrical Engineering, Xiamen University, Xiamen, 361005, China. E-mail addresses: [email protected], [email protected] (W. Zhou). http://dx.doi.org/10.1016/j.sna.2016.04.044 0924-4247/© 2016 Elsevier B.V. All rights reserved.
backing structure in bioelectrodes are also used to reduce the problem of motion interference during the measurement process. Until now, the developed biomedical electrodes are mainly traditional silver/silver chloride electrodes, microneedle electrodes, textile flexible electrodes, flexible substrate electrodes, foam structure electrodes, and insulation dry electrodes [6–8]. In recent years, biomedical electrodes technology has attracted increasing attention from researchers. The design and manufacturing technology of biomedical electrodes has made rapid progress. Nowadays, biomedical electrodes can be divided into two types: dry and wet bioelectrodes. Compared with wet bioelectrodes, i.e., Ag/AgCl electrodes, dry bioelectrodes are attractive because there is no need for an electrolytic gel for the bioelectricity impedance technique, which can effectively overcome the limitations of wet electrode technology and show excellent signal collection performance [3,4]. In particular, a dry bioelectrode with surface microstructure arrays has been developed for its good electrodeskin contact interface to prevent high impedance. Thus, a dry bioelectrode can be used in the long-term recording process without conductive gel and skin preparation, which also has several obvious advantages such as lower impedance variations, smaller electrochemical noise, better stability, and convenient use. Several fabrication methods for bioelectrodes, micromachining such as wet etching processing [9–12], reactive ion etching [13,14], the casting
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Fig. 1. Structural design of dry bioelectrode with a novel construction: 1-Back side of electrode back; 2-Front side of electrode; 3-Shielding wire; 4-Conductive silver glue; 5-Metal electrode core with surface microstructures arrays; 6-Foam backing material.
method [15], and 3D printing [16] have been developed to measure the physiological signal. Up to now, much research work about the fabrication and performance of bioelectrodes have been reported. The wet-etching technique was first proposed to fabricate bioelectrodes with surface micro needle structures [9–12]. Later, Griss et al. [13,14] developed the deep reactive ion etching method to fabricate spike electrodes using silicon that was subsequently covered with a Ag/AgCl double layer to reduce electrode interface noise. Ng et al. [15] developed the vacuum casting method to fabricate microspike dry EEG electrodes, characterized the sensing performance in terms of the impedance level and stability, and showed a much higher efficiency in EEG measurements. Salvo et al. [16] develop the 3D printing technology to fabricate electrodes with 180 conical needles. The ECG and EEG were then measured compared with planar wet Ag/AgCl electrodes. Wang et al. [17] developed two combined methods to fabricate silicon-based dry electrodes for measuring biological signals. Lopez-Gordo et al. [18] review current approaches to develop dry EEG electrodes with nano-, micro, and millineedles for clinical and other applications, including information about measurement methods and evaluation reports. Recently, Liu et al. [19] proposed a micromolding technology to fabricate a three-dimensional microneedle electrode arrays. Electrical impedance and polarization voltage experimental results showed that the electrode chips exhibited great electric characteristics. In the present study, laser micromilling technology was used to fabricate a metal dry electrode with a surface microstructure array, which was then treated with stearic acid ethanol solution. The contact angle of the bioelectrode was experimentally investigated by the varying the laser processing parameters, the spacing of the microstructure, and working media.
Fig. 2. Optical image for laser micromilling setup of dry bioelectrode with surface microstructures arrays. Table 1 Specifications of characteristic parameters of the used fiber laser system. Characteristic
Parameter range Process conditions
Unit
Wavelength Nominal average output power Pulse duration Repetition rate Beam quality (M2 ) Incident beam diameter Focused diameter The number of scans Laser output power Scanning speed
1055–1070 19–21 90–120 20–200 <1.1 6–9 24.3–37.3 0–50 0–0 0–1000
nm W ns kHz
1064 20 100 20 1 7 31.5 5, 10, 20, 50 14, 16, 18, 20 250, 500, 750, 1000
mm m / W mm/s
The specifications of the characteristic parameters of the fiber laser system are given in Table 1. Moreover, a linear motor was used to move the laser head in the z-direction so that the laser beam could be focused on the surface of the electrode sample materials by setting the focal length to 163 mm. To fabricate the bioelectrode with surface microstructures arrays, the cross machining route and the loop multiple-pass reciprocating scanning strategy were predetermined in this study [20–22]. During the laser micromilling process, the surface material can be removed layer by layer through repeating the scan sequence. After the laser micromilling process, the dry bioelectrodes were cleaned with diluted hydrochloric acid and ethanol in an ultrasonic bath for 5 min to remove surface oxides and any organic substance. Then the dry bioelectrode was completely impregnated in 2 g/L stearic acid ethanol solution for 2 h at 40 ◦ C. Finally, the dry bioelectrode was clean with deionized water and dried in a baking box. The surface morphology of the dry bioelectrodes was observed through a scanning electron microscope (SEM) (Hitachi SU70, Japan) with energy-dispersive X-ray measurements (EDX). .
2. Experimental procedure 2.2. Contact angle test setup 2.1. Design and fabrication of dry bioelectrode Fig. 1 shows the schematic diagram of structural design of dry bioelectrode with a novel construction. The novel designed bioelectrode is composed of metal electrode core, foam backing material, conductive silver glue and shielding wire. Optical image for laser micromilling setup of dry bioelectrode is shown in Fig. 2. In this study, a prototype pulsed fiber laser (IPG, No: YLP-1-100-20-20CN, Germany) was used as the fabrication laser. The laser was set to produce 100-ns pulses with a 1064-nm central emission wavelength at a repetition rate of 20 kHz. The nominal output power of the laser was 20 W, which was sufficiently higher than the ablation threshold of the bioelectrode sample materials (red copper).
The contact angle () measurement was carried out with an automatic contact angle meter (Model: DCAT21, Dataphy Instrument Co., Ltd, Germany). The test system was composed of a syringe pump, a CCD camera, a light source, a sliding table, and a computer analyzing system, as shown in Fig. 3. In this study, the static contact angles were actually obtained by sessile drop measurements, where a drop was deposited on the surface, and the value was obtained by a goniometer [23]. To measure the contact angle, a syringe pump was first used to inject a controlled volume of ∼3 L droplet of the working media on the sample surface through the syringe. After the droplet remained on the sample surface for about 30 s, images of the droplet were taken using the CCD camera.
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surements were estimated to be 1◦ . All the measurements were performed at room temperature (approximately 25 ◦ C). 3. Results and discussion 3.1. Surface microstructure of the bioelectrode
Fig. 3. Optical image of contact angle test setup for bioelectrode.
Finally, three images were analyzed using the supplied software to determine the contact angle. For each sample, the contact angle measurement was repeated on five randomly selected areas on the surface of the bioelectrode [24]. The average value of their contact angles was taken as the static contact angle of the samples. The standard deviation and the average error for contact angle mea-
In previous study, different surface micro/nanostructures have been fabricated using laser micromachining process [25,26]. The laser micromachining process can be simply defined as a material removal process by heating the surface material with a laser beam, which includes four processing stages of surface material heating, surface material melting, material evaporation, and ionization as well as evaporation spray and fluid ejection [27,28]. When the fiber laser emitted a collimated beam of white light to continuously heat the material surface, surface thermal energy accumulated quickly, resulting in a significant increase in the temperature of the material surface. After the temperature of the material surface was increased to slightly lower than the melting point of the material, the liquid phase damaged the material surface. Subsequently, with continuous irradiation of the laser beam, metal steam began to partially ionize to form the low-density plasma, which was helpful to further absorb laser energy. Because of the pressure of the vaporizing expansion and plasma, the surface liquid material gradually evaporated to induce the ejection phenomenon to produce many grooves.
Fig. 4. Typical images of the microstructure array on the surface of the dry bioelectrode:(a) Optical image; (b) SEM image; (c) Cross-section view; (d) Top view.
Fig. 5. Contact angle on the bioelectrode surface under different conditions: (a) untreated, (b) after one month in air, and (3) treated with stearic acid ethanol solution.
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7
7 C O
Weight (%)
6
6
5
5
4
4
3
3
2
2
1
1
0
a
b
c
0
Condition Fig. 6. The proportion of carbon and oxygen by weight of the bioelectrodes: (a) untreated, (b) after one month in air, and (3) treated with stearic acid ethanol solution.
In particular, a recasting phenomenon readily occurred to produce the recast layer structure to deposit the unscanned area between the process spacings. Thus, a small quantity of surface material was removed from the laser-scanned area, and the recast layer started to form around the unscanned area. When the number of scans was gradually increased, more recast layer was continually deposited and solidified on the unscanned area. Ultimately, the recast layer was deposited and combined on the unscanned area to form a microstructure array with different geometrical dimensions on the surface of the sample material. Typical images of the microstructure array on the surface of the dry bioelectrode are shown in Fig. 4 The surface microstructure array of the dry bioelectrode was successfully fabricated using the recast layer in the laser micromilling process.
3.2. Contact angle on the bioelectrode surface under different conditions Fig. 5 shows the contact angle on the electrode surface under different conditions. The bioelectrode with 0.1 mm spacing of surface microstructure was fabricated with a scanning speed of 500 mm/s, a laser power of 20 W, and 20 scans. For the untreated bioelectrode, water drops spread out immediately on the bioelectrode surface, so the contact angle was 0◦ . When the bioelectrode was placed in the air for one month, the contact angle was 140◦ . This indicated that the bioelectrode showed hydrophobic properties. Several previous studies also shows that micro/nano-structured metal surface become superhydrophobic from initial hydrophilic in air over time [29–31]. In particular, the contact angle was 151◦ when the bioelectrode was treated with stearic acid ethanol solution. Thus, a bioelectrode with superhydrophobicity could be obtained. Fig. 6 shows that the proportion of carbon and oxygen by weight changed significantly. The contact angle increased with the carbon weight. However, the contact angle did not change with the oxygen weight, because the C C bond influences the surface polarity. A higher C C value indicates a more nonpolar surface, which tends to hydrophobicity [32,33]. Compared with a copper sheet treated with stearic acid ethanol solution ( = 103◦ ), the contact angle of the bioelectrode with a surface microstructure array fabricated using the laser micromilling technique increased greatly. These results indicated that the surface microstructure array could change the state of contact with water droplets. A bioelectrode with surface hydrophobic properties could be obtained.
Fig. 7. Contact angle of bioelectrodes fabricated with different laser process parameters.
3.3. Effect of laser processing parameters on the contact angle Fig. 7 shows the contact angle of a bioelectrode fabricated with different laser processing parameters. With a laser output powers of 14W–20W, the contact angle increased with the increasing laser output power. After the laser output power increased to above 18 W, the contact angle reached a stable value of around 150◦ . When the scanning times were increased, with values of 5, 10, 20, and 50 times, the contact angle increased, as shown in Fig. 7b. When the scanning speed was less than 750 mm/s, the contact angle changed little. However, a slightly decreasing trend was observed when the scanning speed increased to 1000 mm/s, as shown in Fig. 7c. These results can be attributed that the surface morphology and size (such as height and spacing) of the microstructure array on the bioelectrode surface changed with different laser processing parameters [21]. Therefore, reasonable laser processing parameters, including
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Fig. 8. SEM image of bioelectrodes with different microstructure spacings.
0.1 mm and 0.2 mm. When the microstructure spacings increased to 0.5 mm, the contact angle deceased to 146◦ . Therefore, we conclude that the microstructure spacing can affect the value of the contact angle. Furthermore, all contact angles of bioelectrodes were above 146◦ , indicating good hydrophobic properties. For the small microstructure spacing, the droplets were lifted from the electrode surface. The contact angle was thus increased. Thus, the developed bioelectrodes with smaller microstructure spacing lift the droplets to exhibit the Cassie state [34,35]. However, with increasing microstructure spacing, the droplets can easily enter the gaps in the microstructure. The contact status also gradually shifted from the Cassie state to the Wenzel state, leading to a smaller contact angle. 3.5. Different working media on the bioelectrode surface Fig. 9. Contact angle of bioelectrodes with different microstructure spacings.
the laser output power, scanning time, and scanning speed, should be selected to fabricate bioelectrodes with much larger contact angles. 3.4. Effect of microstructure spacing on the contact angle Figs. 8 and 9 show SEM images and contact angles of bioelectrodes with different microstructure spacings. The contact angle increased with a decrease in the microstructure spacing. The contact angle can be above 150◦ when the microstructure spacings are
Because dry bioelectrodes are in direct contact with the related working medium during long-term measurement processes, it is important to study the contact angle of the electrode surface with different working media [16]. In particular, sweat is easily produced on the skin because of the body’s metabolic activity, which seriously affects the measurement results. Therefore, deionized water, 3 g/L saline water (similar to the sweat of sodium chloride concentration), and glycerin were selected as working media for the contact angle measurement. The results show that the contact angle of deionized water and saline water on the electrode surface could reach above 150◦ , indicating superhydrophobic properties. The contact angle for glycerin was 142◦ . In general, the dry electrodes showed good hydrophobicity and oleophobic properties, as
Fig. 10. Deionized water, saline water and glycerin droplets on surface of bioelectrode: (a) Optical images; (b) Contact angle value.
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shown in Fig. 10. When the dry electrode is in contact with deionized water, saline water, and glycerin, a larger contact angle can be obtained to maintain good contact conditions between the electrode and skin. Thus, the stability of the measurement process with a dry bioelectrode can be improved. 4. Conclusions The laser micromilling technology was used to fabricate dry bioelectrodes with a surface microstructure array, which were then treated with stearic acid ethanol solution. The contact angle of a dry bioelectrode could reach 151◦ , indicating superhydrophobic properties. After discussing the formation of the dry bioelectrode with a surface microstructure array, the contact angle on the electrode surface was analyzed by changing the parameters, such as through laser processing parameters, spacing of the microstructure, and the working media. The laser output power and the number of scans exhibited a great influence on the contact angle, but the scanning speed did not have much influence. The dry bioelectrodes with microstructure spacings of 0.1 mm and 0.2 mm presented good hydrophobic properties. In addition, the contact angles on the bioelectrode with deionized water and saline water as the working media were greater than 150◦ . Our developed metal dry bioelectrode presented a good hydrophobic properties to help realize stable measurements over a long time and have wide application prospects in the detection and measurement of biological signals. Acknowledgments This work was supported by the National Natural Science Foundation of China (Project No. 51475397) and the Fundamental Research Funds for Central Universities, Xiamen University (Project No. 20720160079). In addition, the supports from Open Fund of Shanghai Key Laboratory of Digital Manufacture for Thin-Walled Structures (No. 2015004) are also acknowledged. References [1] Y.M. Chi, T.P. Jung, G. Cauwenberghs, Dry-contact and noncontact biopotential electrodes: methodological Review, IEEE Rev. Biomed. Eng. 3 (2010) 106–119. [2] E. Spinelli, M. Haberman, Insulating electrodes: a review on biopotential front ends for dielectric skin-electrode interfaces, Physiol. Meas. 31 (2010) s183–s198. [3] M. Bodenstein, M. David, K. Markstaller, Principles of electrical impedance tomography and its clinical application, Crit. Care Med. 37 (2009) 713–724. [4] A. Searle, L. Kirkup, A direct comparison of wet, dry and insulating bioelectric recording electrodes, Physiol. Meas. 21 (2000) 271–283. [5] J.Y. Baeka, J.H. An, J.M. Choi, Flexible polymeric dry electrodes for the long-term monitoring of ECG, Sens. Actuators A 143 (2008) 423–429. [6] A. Gruetzmann, S. Hansen, J. Muller, Novel dry electrodes for ECG monitoring, Physiol. Meas. 28 (2007) 1375–1390. [7] C.T. Lin, L.D. Liao, Y.H. Liu, Novel dry polymer foam electrodes for long-term EEG measurement, IEEE T. Bio-Med. Eng. 58 (2011) 1200–1207. [8] W. Zhou, W. Liu, Q.F. Qiu, Development, fabrication, and applications of biomedical electrodes (in Chinese), Chin. Sci. Bull. 60 (2015) 1352–1360. [9] P.K. Campbell, K.E. Jones, R.J. Huber, K.W. Horch, R.A. Normann, A silicon-based three-dimensional neural interface-manufacturing processes for an intracortical electrode array, IEEE Trans. Biomed. Eng. 38 (1991) 758–768. [10] Q. Bai, K.D. Wise, D.J. Anderson, A high-yield microassembly structure for three-dimensional microelectrode arrays, IEEE Trans. Biomed. Eng. 47 (2000) 281–289. [11] C.O. Mahonya, F. Pini, A. Blakea, C. Webstera, J.O. Briena, K.G. McCarthy, Microneedle-based electrodes with integrated through-silicon via for biopotential recording, Sens. Actuators A 186 (2012) 130–136. [12] N. Wilke, C. Hibert, J.O. Brien, A. Morrissey, Silicon microneedle electrode array with temperature monitoring for electroporation, Sens. Actuators A 123–124 (2005) 319–325. [13] P. Griss, P. Enoksson, H.K. Tolvanen-Laakso, P. Merilainen, S. Ollmar, G. Stemme, Micromachined electrodes for biopotential measurements, J. Microelectromech. S. 10 (2001) 10–16. [14] P. Griss, H.K. Tolvanen-Laakso, P. Merilainen, G. Stemme, Characterization of micromachined spiked biopotential electrodes, IEEE T. Bio-Med. Eng. 49 (2002) 597–604.
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Biographies Wei Zhou received his B.S. degree in mechanical engineering from Xiangtan University, Xiangtan, China, in 2005. In 2010, he received his Ph.D. degree in mechanical engineering from South China University of Technology, Guangzhou, China. From 2010–2012, he was a postdoc researcher at Sun Yat-sen University. Now he is an associate professor at Xiamen University. His current research interests focus on design, fabrication and performance evaluation of biomedical device. Wei liu starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2014. Her research interests are in the area of measure methods of different bioelectrodes. Shaoyu Liu starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University 2014. His research interests are in microfabrication technology of biomedical device. Guobiao Zhang starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2013. Her research interests are in the area of modeling and simulation of microfabrication process of bioelectrode. Zhijia Shen starts to work toward the Ph.D. degree in mechanical engineering at Xiamen University since 2013. His research interests are in novel measurement technology for ECG and EMG.