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Low melting point metal-based flexible 3D biomedical microelectrode array by phase transition method
Materials Science & Engineering C 99 (2019) 735–739
Contents lists available at ScienceDirect
Materials Science & Engineering C journal homepage: www.elsevier.com/locate/msec
Low melting point metal-based flexible 3D biomedical microelectrode array by phase transition method Shengxin Guoa, Rongzan Lina, Lei Wangb, Shinying Laua, Qian Wangb, Ran Liua,
T
⁎
a
Department of Biomedical Engineering, School of Medicine, Tsinghua University, Beijing 100084, China Beijing Key Lab of Cryo-Biomedical Engineering and Key Lab of Cryogenics, Technical Institute of Physics and Chemistry, Chinese Academy of Sciences, Beijing 100190, China
b
A R T I C LE I N FO
A B S T R A C T
Keywords: Flexible Microelectrode Low melting point metal Liquid metal Fast fabrication Phase transition 3D printing
In this study, we present a dimension-controllable 3D biomedical microelectrode based on low melting point metals (Bi/In/Sn/Zn alloy) applied using the phase transition method. We have established a process, in which the liquid metal is pumped through a syringe needle of the dispensing system to form a needle shape after cooling at room temperature. PDMS (polydimethylsiloxane) was chosen as the substrate of the electrode as it is amenable to micro-molding and has excellent flexibility. Several key factors, including lifting velocity of the syringe needle and sample temperature were examined as to how they would affect the height, width and depthwidth ratio of the electrode, to realize size control of the electrode. Afterwards, the skin-electrode impedance was tested and the results were compared with those of an Ag/AgCl (wet) electrode. The impedance at 10 Hz is 2.357 ± 0.198 MΩ for the 3D microelectrode. From data, the impedance of 3D microelectrode is found to be at the same level as the Ag/AgCl electrode at the frequency of 10 Hz. By increasing the size of the array, the impedance of the low melting point metal electrode and the wet electrode converge. The resistance of the electrode was also measured to describe its stretchability. The electrode can be stretched to a maximum of 42% before it becomes non-conducting. In addition to acquisition of bio-electric signals, our method has strong prospects in the field of bio-sensing.
1. Introduction Biopotential signals, such as ECG, EEG and EMG provide an effective way to diagnose diseases and monitor the human body. The choice of electrode plays an important role in acquiring high quality bioelectric signals for this purpose. The conventional silver/silver chloride (Ag/AgCl) wet electrode is the most commonly used, and is used for acquiring bio-electric signals due to its efficient electrical coupling, which is provided by the conductive gel [1]. However, it may cause skin irritation, and the drying of the gel may lead to degradation of electrical performance during long-term monitoring [2]. Recently, wet electrodes have been replaced by dry electrodes, such as microneedles, contact probes, capacitive disks, conductive composites and nanowire electrodes [3–9]. Of the alternative methods listed above, the 3D electrode is superior to the planar electrodes with its superior transdermal characteristics, which leads to low skin-electrode impedance [10–13]. However, the fabrication process of 3D electrode is extremely complicated, and includes material deposition, photo-etching and ion etching [14–16].
⁎
A traditional 3D electrode typically consists of a rigid metal or silicon substrate with thickness in the range of several millimeters. The rigid interface between the skin and electrode does not offer reliable and comfortable use by human users, which constrains its wider use in long-term bio-electric signals monitoring [17]. Accordingly, research now aims to achieve flexibility in electrodes. Micro-Electro-Mechanical Systems (MEMS) technology and mixed materials are widely applied in the manufacture of flexible electrodes, however, most fabrication methods have several disadvantages such as complexity, low yield, etc. [18]. To achieve high flexibility and avoid a complicated fabrication process, we propose the use of a low melting point metal alloy for rapid fabrication using the phase transition method. The distinctive metal alloy can easily shift between liquid and solid phase by a change in the temperature. Therefore, this metal alloy and the related processes have found application in many fields. There have been many state-of-the-art innovations in rapid manufacture and in 3D printing using this metal alloy, by virtue of its good biocompatibility and excellent fluidity beyond its melting point. For these reasons, it has been investigated as a
Corresponding author. E-mail address: [email protected] (R. Liu).
https://doi.org/10.1016/j.msec.2019.02.015 Received 29 November 2017; Received in revised form 4 February 2019; Accepted 4 February 2019 Available online 05 February 2019 0928-4931/ © 2019 Elsevier B.V. All rights reserved.
Materials Science & Engineering C 99 (2019) 735–739
S. Guo, et al.
Fig. 1. Formation of droplet-shaped metal with sharp needle. (A) The low melting point metal was pumped through the syringe needle. (B) The syringe needle was lifted up after the alloy had adequate contact area with the flexible substrate. (C) The low melting point metal retains its droplet shape with sharp needle after cooling down at room temperature. Schematic illustration of the 3D flexible microelectrode fabrication process. (D) The dispensing system can form any size array of the droplet-shaped metal with sharp needles. (E) The Ga/In alloy in liquid state was sprayed on the surface of the electrode array to make it conductive. (F) Flexible flat cable (FFC) was placed on the surface of the conductive layer. (G) PDMS pre-polymer was coated on the electrode and conductive layer to achieve isolation.
point (57.5 °C), which makes it easy to shape. All the beakers were cleaned with de-ionized water and metals of 99.99% purity were placed in the beaker. The beakers were then heated for 5 h at 245 °C in an electric vacuum drying oven until the metals had molten and fused completely followed by stirring with a magnetic stirrer at 70 °C for 4 h to further increase the homogeneity of the components. After the mixing process, the alloy was added into the injection syringe and stored at room temperature. The conductive layer was made of gallium (Ga) and indium (In) printable metals with the proportion of 75.5% and 24.5% by weight, respectively. The gallium and indium metal of purity 99.99% were added into a prepared baker and were heated at 100 °C until they were molten and fused completely. Then, 8 ml of 30% NaOH solution was added to the alloy to prevent the formation of natural oxidation as the melting point of Ga/In alloy is as low as 15 °C, which ensures it remains in a liquid state at room temperature. Experimental devices: In the fabrication process, a dispenser (TYSR200D, Beijing Tinyo Electronics Co., Ltd., China) was used to dispense Bi/In/Sn/Zn alloy on the substrate through a syringe needle (TE732050, 0.9 mm inner diameter); the parameters including dispensing time, trail time, lifting velocity of the syringe needle, the distance between the syringe and the substrate, and the shape of the dot array
suitable raw material for bone cement and printing ink in direct writing electronics and 3D printing technology [19–21]. In this study, an innovative approach was adopted to further improve the techniques for fabricating flexible 3D microelectrodes. The microelectrodes, with sharp tips and specific height, were fabricated from a Bi/In/Sn/Zn alloy using a dispensing system. The dimension of the microelectrodes can be controlled by varying the temperature of the sample material as determined from the system parameters. The experiments revealed several pivotal fundamental electrical properties of the 3D microelectrodes. The developed electrodes have potential application in the monitoring of bio-electric signals by just applying physical stress to the substrate. To summarize, the fabrication of dimension-controllable 3D microelectrodes on an elastic substrate using phase transition is both simple and quick.
2. Methods Preparation of the materials: The electrode is made of an alloy of bismuth (Bi), indium (In), tin (Sn) and zinc (Zn) metals with proportion of 35%, 48.6%, 16% and 0.4% by weight, respectively. This alloy has been chosen because of its high electrical conductivity and low melting 736
Materials Science & Engineering C 99 (2019) 735–739
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The greatest merit of our method is that the size of our 3D microelectrode can be easily controlled by applying its phase transition property during the sampling process to adapt it to various applications. We set the lifting velocity of the syringe needle as 10–120 mm/s with an interval of 10 mm/s and the sample temperature as 75 °C and 90 °C, respectively, while other experiment parameters were kept the same. For each group in Fig. 3A, B and C, we produced 6 samples. From the results, we can preliminarily conclude that as the velocity of the syringe increases, the height of the microelectrode declines, while higher sampling temperature tends to lead to greater electrode height (Fig. 3A). Notably, the width of the electrode remains nearly the same at different lifting speeds and sample temperatures (Fig. 3B). Therefore, the depth-width ratio of the electrode can be easily controlled just by changing the parameters, which include the lifting velocity of the syringe needle and initial sample temperature (Fig. 3C). In acquiring bio-electric signals, the skin-electrode impedance should be kept low and stable. Accordingly, the impedance of both Ag/ AgCl wet electrode and low melting point metal electrode were measured through an experiment, to reveal the impedance characteristic of our 3D microelectrode. We used the 4 × 4 3D microelectrode array with PDMS insulation layer. The height of electrode was 269 μm and the width was 96 μm. The impedance of the low melting point metal and the Ag/AgCl electrode were measured by an electrochemical workstation, respectively. We supposed that the electrode can be used in acquiring EEG and ECG. Usually in human use, the frequency of EEG is 0–20 Hz while the frequency of ECG is 0.5–100 Hz. Hence, we choose 10 Hz as the frequency to compare the impedance. The experiment was repeated 3 times. The impedance at 10 Hz is 2.357 ± 0.198 MΩ for the 3D microelectrode while the value is 0.973 ± 0.157 MΩ for the Ag/ AgCl electrode (Fig. 3D). From the data, the impedance of the 3D microelectrode is at the same level with the wet electrode at a frequency of 10 Hz. By increasing the size of the array, the impedance of the low melting point metal 3D electrode and the wet electrode converges. We also measured the stretching range of the low melting point electrode to ensure its stretchability in bio-electrical signal monitoring. We used the 4 × 4 low melting point electrode array with the PDMS substrate. The height of the electrode was 285 μm and the width was 99 μm. We stretched the PDMS substrate of the electrode by the edge in single axis and measured its resistance afterwards. The initial length of the substrate was 1.4 cm (Fig. 3E). It was not conductive when stretched to more than 2 cm, which indicated that the electrode can be stretched by a maximum of 42%. In addition, multi-channel microelectrode arrays can also be developed by applying this method. First, the liquid metal conductor (Ga/ In alloy) was printed by a liquid metal printer on the PVC substrate, then PDMS pre-polymer was coated on the circuit. After being cured at 80 °C in the oven for 1 h, the circuit on the PVC was transferred to the PDMS substrate (Fig. 4A). Afterwards, we aligned the syringe with the connection point in turn and dispensed the microelectrode array (Fig. 4B). Finally, FFC was placed on the connector pad and the PDMS pre-polymer was coated on the liquid metal cured at room temperature for 48 h (Fig. 4C–D). By this approach, the multi-channel flexible electrode was successfully fabricated (Fig. 4E).
were controlled by a computer. Since the alloy is in solid state at room temperature, we fitted an aluminum alloy cylinder heated by constantan wire of resistance 62 Ώ/m outside the plastic syringe to melt the alloy. By supplying electrical power to the constantan resistance wire, the metal was heated from 75 °C to 90 °C, which was monitored by a data collecting instrument (Agilent 34972A, Agilent Technologies Inc., USA). A nitrogen cylinder was used to supply a constant pressure, which was regulated by a solenoid valve on the syringe. In addition, a paint gun and air bottle were used to spray the Ga/In alloy on the surface of the electrode array. Impedance experiment: The impedance data was collected by an electrochemical workstation (Princeton PARSTAT 4000, USA) to assess the electrical properties of the microelectrode. We measured the impedance of the microelectrode and the wet gel electrode. Two shielded copper wires linked with Ga/In alloy conductive film by a clamp, were used to provide reliable connection between the electrode and the equipment. Two microelectrodes were placed on the subject's forearm with 5 cm between them. 10 mV root mean square (RMS) sinusoidal voltage with a frequency range from 1000 Hz to 10 Hz was provided for impedance measurement. Impedance of the Ag/AgCl wet electrode with 0.8 cm diameter was also measured in the same conditions, for comparison. 3. Results and discussion A four-element alloy Bi35In48.6Sn15.9Zn0.4 (melting point: 57.5 °C) was used as the raw material for the electrode. The melting enthalpy and the specific heat capacity of Bi35In48.6Sn15.9Zn0.4 are 28.94 J/g and 0.262 J/(g*°C), respectively, which are much lower than those of commonly used metals. These properties enable its easy liquid-solid phase transition during the dispensing process. After being heated to liquid state, the alloy was pumped through the syringe needle, which was then lifted after the liquid alloy shared adequate contact area with the flexible substrate (Fig. 1A–B). At room temperature, the low melting point metal alloy cools down immediately in to a droplet shape with sharp needle (Fig. 1C). By applying this approach, a low melting point metal can be used to rapidly fabricate flexible 3D microelectrode arrays. The process begins with the PDMS being mixed with a curing agent in the ratio of 10:1, after which the mixture was degassed in a vacuum chamber for 30 min. The pre-polymer was heated at 80 °C for 1 h to solidify and form the substrate of the electrode. Then, the dispensing system was prepared to an adequate temperature to melt the alloy. By setting parameters in the computer program controlling the dispensing process, the droplet-shaped metal with sharp needle can be dispensed to form any size array (Fig. 1D). To make the separated electrodes conductive, Ga/In alloy in liquid state at room temperature was sprayed on the surface of the electrode array (Fig. 1E). Finally, a flexible flat cable (FFC) was placed on the surface of the conductive layer and PDMS prepolymer coated on it followed by curing at room temperature for 48 h, to achieve electrical isolation (Fig. 1F–G). The 4 × 4 microelectrode array before and after sealing was observed through a camera lens and CCD (IMAVISION MER-U3x, China), respectively (Fig. 2A–B). The PDMS sealing layer fixed the electrode on the PDMS substrate and prevented direct contact between the Ga/In alloy and skin. The 6 × 6 microelectrode array was successfully fabricated (Fig. 2C) within 5 min, which suggests fast fabrication and easy shape control of the array using our approach. Generally, the needles of the 3D microelectrode should pierce the stratum corneum (40–100 μm) to lower the skin-electrode impedance but avoid reaching the epidermal layer (200–400 μm), which would cause pain, thus, the height of the 3D microelectrode should be in the range of 100–400 μm. The specific dimension of the electrode was measured by scanning electron microscope (Fig. 2D–F). The average length of the needles was found to be 269.985 ± 25.569 μm, which satisfies the transdermal requirement.
4. Conclusions In summary, we succeeded in directly fabricating a 3D microelectrode using phase transition method. This technology enables the rapid formation of a flexible 3D microelectrode array because of its transition mechanics. The dimension of the electrode and the size of the array can be easily controlled by just setting the system parameters including lifting velocity of the syringe needle and sample temperature. By increasing the sample temperature, the height of the electrode can reach up to 1 mm. It may also be noted that the height of the sharp needle of the electrode can be adjusted to penetrate just up to the stratum corneum, which decreases the skin-electrode impedance while protecting 737
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Fig. 2. The microelectrode array as observed through a camera lens and CCD. (A) The unsealed 4 × 4 microelectrode array. (B) The 4 × 4 3D microelectrode array with PDMS insulation layer. (C) A 6 × 6 microelectrode array as observed through a scanning electron microscope (SEM). (D) The microelectrode array. (E) The single needle of the electrode. (F) The pin point of the electrode.
Fig. 3. (A) The relationship between the height of the electrode and lifting velocity at different sample temperatures. (B) The relationship between the width of the electrode and lifting velocity at different sample temperatures. (C) The relationship between the depth-width ratio of the electrode and lifting velocity at different sample temperatures. (D) The skin-electrode impedance of the low melting metal electrode and Ag/AgCl electrode at frequency ranging from 10 Hz to 1000 Hz, respectively. (E) The relationship between the resistance of the low melting point metal electrode and the length of the electrode pad.
practical issues to be solved in the near future.
the patient from feeling any pain. Apart from applying in acquiring bioelectric signals, the 3D microelectrode has strong potential in bio-sensing and preparation of electrode templates because of its flexibility and fast fabrication characteristic. Our subsequent research will aim to acquire ECG and EEG signals using our electrodes. As a completely new fabrication method, our research raises important fundamental and
Acknowledgements This research was supported by the National Natural Science Foundation of China under Grant No. 81471749, the National Key R&D 738
Materials Science & Engineering C 99 (2019) 735–739
S. Guo, et al.
Fig. 4. (A)–(D) Schematic illustration of the fabrication process of multi-channel flexible microelectrode array device. (E) Image of multi-channel flexible microelectrode.
Program of China under Grant No. 2017YFC0803507, the Tsinghua University Initiative Scientific Research Program under Grant No. 2014Z01001, and the Foundation of Beijing Laboratory in Biomedical Technology and Instruments. We express gratitude to Capital Bio Corporation Micro Systems Engineering Lab for their help.
(2014). [9] A.R. Mota, L. Duarte, D. Rodrigues, A.C. Martins, A.V. Machado, F. Vaz, P. Fiedler, J. Haueisen, J.M. Nóbrega, C. Fonseca, Sensors Actuators A Phys. 199 (2013). [10] G. Stavrinidis, K. Michelakis, V. Kontomitrou, G. Giannakakis, M. Sevrisarianos, G. Sevrisarianos, N. Chaniotakis, Y. Alifragis, G. Konstantinidis, Microelectron. Eng. 159 (2016). [11] W.C. Ng, H.L. Seet, K.S. Lee, N. Ning, W.X. Tai, M. Sutedja, J.Y.H. Fuh, X.P. Li, Mater. Process. Technol. 209 (2009) 9. [12] G. Ruffini, S. Dunne, L. Fuentemilla, C. Grau, E. Farrés, J. Marco-Pallarés, P.C.P. Watts, S.R.P. Silva, Sensors Actuators A Phys. 144 (2) (2008). [13] M. Matteucci, R. Carabalona, M. Casella, E. Di Fabrizio, F. Gramatica, M. Di Rienzo, E. Snidero, L. Gavioli, M. Sancrotti, Microelectron. Eng. (2007) 84. [14] H. Zhang, W. Pei, Y. Chen, X. Guo, X. Wu, X. Yang, H. Chen, IEEE Trans. Biomed. Eng. 63 (2016) 6. [15] L.F. Wang, J.Q. Liu, X.X. Yan, B. Yang, C.S. Yang, Microsyst. Technol. 19 (2013) 2. [16] D.H. Ren, M.Y. Cui, Y.Q. Xia, Z. You, Prog. Biochem. Biophys. 39 (2012) 10. [17] R. Liu, X. Yang, C. Jin, J. Fu, W. Chen, J. Liu, Appl. Phys. Lett. 103 (2013) 19. [18] R. Wang, W. Zhao, W. Wang, Z. Li, J. Microelectromech. Syst. 21 (2012) 5. [19] L. Yi, C. Jin, L. Wang, J. Liu, Biomaterials 35 (2014) 37. [20] L. Wang, J. Liu, Sci. China Technol. Sci. 57 (2014) 11. [21] L. Wang, J. Liu, Sci. China Technol. Sci. 57 (2014) 9.
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Contents lists available at ScienceDirect
Materials Science & Engineering C journal homepage: www.elsevier.com/locate/msec
Low melting point metal-based flexible 3D biomedical microelectrode array by phase transition method Shengxin Guoa, Rongzan Lina, Lei Wangb, Shinying Laua, Qian Wangb, Ran Liua,
T
⁎
a
Department of Biomedical Engineering, School of Medicine, Tsinghua University, Beijing 100084, China Beijing Key Lab of Cryo-Biomedical Engineering and Key Lab of Cryogenics, Technical Institute of Physics and Chemistry, Chinese Academy of Sciences, Beijing 100190, China
b
A R T I C LE I N FO
A B S T R A C T
Keywords: Flexible Microelectrode Low melting point metal Liquid metal Fast fabrication Phase transition 3D printing
In this study, we present a dimension-controllable 3D biomedical microelectrode based on low melting point metals (Bi/In/Sn/Zn alloy) applied using the phase transition method. We have established a process, in which the liquid metal is pumped through a syringe needle of the dispensing system to form a needle shape after cooling at room temperature. PDMS (polydimethylsiloxane) was chosen as the substrate of the electrode as it is amenable to micro-molding and has excellent flexibility. Several key factors, including lifting velocity of the syringe needle and sample temperature were examined as to how they would affect the height, width and depthwidth ratio of the electrode, to realize size control of the electrode. Afterwards, the skin-electrode impedance was tested and the results were compared with those of an Ag/AgCl (wet) electrode. The impedance at 10 Hz is 2.357 ± 0.198 MΩ for the 3D microelectrode. From data, the impedance of 3D microelectrode is found to be at the same level as the Ag/AgCl electrode at the frequency of 10 Hz. By increasing the size of the array, the impedance of the low melting point metal electrode and the wet electrode converge. The resistance of the electrode was also measured to describe its stretchability. The electrode can be stretched to a maximum of 42% before it becomes non-conducting. In addition to acquisition of bio-electric signals, our method has strong prospects in the field of bio-sensing.
1. Introduction Biopotential signals, such as ECG, EEG and EMG provide an effective way to diagnose diseases and monitor the human body. The choice of electrode plays an important role in acquiring high quality bioelectric signals for this purpose. The conventional silver/silver chloride (Ag/AgCl) wet electrode is the most commonly used, and is used for acquiring bio-electric signals due to its efficient electrical coupling, which is provided by the conductive gel [1]. However, it may cause skin irritation, and the drying of the gel may lead to degradation of electrical performance during long-term monitoring [2]. Recently, wet electrodes have been replaced by dry electrodes, such as microneedles, contact probes, capacitive disks, conductive composites and nanowire electrodes [3–9]. Of the alternative methods listed above, the 3D electrode is superior to the planar electrodes with its superior transdermal characteristics, which leads to low skin-electrode impedance [10–13]. However, the fabrication process of 3D electrode is extremely complicated, and includes material deposition, photo-etching and ion etching [14–16].
⁎
A traditional 3D electrode typically consists of a rigid metal or silicon substrate with thickness in the range of several millimeters. The rigid interface between the skin and electrode does not offer reliable and comfortable use by human users, which constrains its wider use in long-term bio-electric signals monitoring [17]. Accordingly, research now aims to achieve flexibility in electrodes. Micro-Electro-Mechanical Systems (MEMS) technology and mixed materials are widely applied in the manufacture of flexible electrodes, however, most fabrication methods have several disadvantages such as complexity, low yield, etc. [18]. To achieve high flexibility and avoid a complicated fabrication process, we propose the use of a low melting point metal alloy for rapid fabrication using the phase transition method. The distinctive metal alloy can easily shift between liquid and solid phase by a change in the temperature. Therefore, this metal alloy and the related processes have found application in many fields. There have been many state-of-the-art innovations in rapid manufacture and in 3D printing using this metal alloy, by virtue of its good biocompatibility and excellent fluidity beyond its melting point. For these reasons, it has been investigated as a
Corresponding author. E-mail address: [email protected] (R. Liu).
https://doi.org/10.1016/j.msec.2019.02.015 Received 29 November 2017; Received in revised form 4 February 2019; Accepted 4 February 2019 Available online 05 February 2019 0928-4931/ © 2019 Elsevier B.V. All rights reserved.
Materials Science & Engineering C 99 (2019) 735–739
S. Guo, et al.
Fig. 1. Formation of droplet-shaped metal with sharp needle. (A) The low melting point metal was pumped through the syringe needle. (B) The syringe needle was lifted up after the alloy had adequate contact area with the flexible substrate. (C) The low melting point metal retains its droplet shape with sharp needle after cooling down at room temperature. Schematic illustration of the 3D flexible microelectrode fabrication process. (D) The dispensing system can form any size array of the droplet-shaped metal with sharp needles. (E) The Ga/In alloy in liquid state was sprayed on the surface of the electrode array to make it conductive. (F) Flexible flat cable (FFC) was placed on the surface of the conductive layer. (G) PDMS pre-polymer was coated on the electrode and conductive layer to achieve isolation.
point (57.5 °C), which makes it easy to shape. All the beakers were cleaned with de-ionized water and metals of 99.99% purity were placed in the beaker. The beakers were then heated for 5 h at 245 °C in an electric vacuum drying oven until the metals had molten and fused completely followed by stirring with a magnetic stirrer at 70 °C for 4 h to further increase the homogeneity of the components. After the mixing process, the alloy was added into the injection syringe and stored at room temperature. The conductive layer was made of gallium (Ga) and indium (In) printable metals with the proportion of 75.5% and 24.5% by weight, respectively. The gallium and indium metal of purity 99.99% were added into a prepared baker and were heated at 100 °C until they were molten and fused completely. Then, 8 ml of 30% NaOH solution was added to the alloy to prevent the formation of natural oxidation as the melting point of Ga/In alloy is as low as 15 °C, which ensures it remains in a liquid state at room temperature. Experimental devices: In the fabrication process, a dispenser (TYSR200D, Beijing Tinyo Electronics Co., Ltd., China) was used to dispense Bi/In/Sn/Zn alloy on the substrate through a syringe needle (TE732050, 0.9 mm inner diameter); the parameters including dispensing time, trail time, lifting velocity of the syringe needle, the distance between the syringe and the substrate, and the shape of the dot array
suitable raw material for bone cement and printing ink in direct writing electronics and 3D printing technology [19–21]. In this study, an innovative approach was adopted to further improve the techniques for fabricating flexible 3D microelectrodes. The microelectrodes, with sharp tips and specific height, were fabricated from a Bi/In/Sn/Zn alloy using a dispensing system. The dimension of the microelectrodes can be controlled by varying the temperature of the sample material as determined from the system parameters. The experiments revealed several pivotal fundamental electrical properties of the 3D microelectrodes. The developed electrodes have potential application in the monitoring of bio-electric signals by just applying physical stress to the substrate. To summarize, the fabrication of dimension-controllable 3D microelectrodes on an elastic substrate using phase transition is both simple and quick.
2. Methods Preparation of the materials: The electrode is made of an alloy of bismuth (Bi), indium (In), tin (Sn) and zinc (Zn) metals with proportion of 35%, 48.6%, 16% and 0.4% by weight, respectively. This alloy has been chosen because of its high electrical conductivity and low melting 736
Materials Science & Engineering C 99 (2019) 735–739
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The greatest merit of our method is that the size of our 3D microelectrode can be easily controlled by applying its phase transition property during the sampling process to adapt it to various applications. We set the lifting velocity of the syringe needle as 10–120 mm/s with an interval of 10 mm/s and the sample temperature as 75 °C and 90 °C, respectively, while other experiment parameters were kept the same. For each group in Fig. 3A, B and C, we produced 6 samples. From the results, we can preliminarily conclude that as the velocity of the syringe increases, the height of the microelectrode declines, while higher sampling temperature tends to lead to greater electrode height (Fig. 3A). Notably, the width of the electrode remains nearly the same at different lifting speeds and sample temperatures (Fig. 3B). Therefore, the depth-width ratio of the electrode can be easily controlled just by changing the parameters, which include the lifting velocity of the syringe needle and initial sample temperature (Fig. 3C). In acquiring bio-electric signals, the skin-electrode impedance should be kept low and stable. Accordingly, the impedance of both Ag/ AgCl wet electrode and low melting point metal electrode were measured through an experiment, to reveal the impedance characteristic of our 3D microelectrode. We used the 4 × 4 3D microelectrode array with PDMS insulation layer. The height of electrode was 269 μm and the width was 96 μm. The impedance of the low melting point metal and the Ag/AgCl electrode were measured by an electrochemical workstation, respectively. We supposed that the electrode can be used in acquiring EEG and ECG. Usually in human use, the frequency of EEG is 0–20 Hz while the frequency of ECG is 0.5–100 Hz. Hence, we choose 10 Hz as the frequency to compare the impedance. The experiment was repeated 3 times. The impedance at 10 Hz is 2.357 ± 0.198 MΩ for the 3D microelectrode while the value is 0.973 ± 0.157 MΩ for the Ag/ AgCl electrode (Fig. 3D). From the data, the impedance of the 3D microelectrode is at the same level with the wet electrode at a frequency of 10 Hz. By increasing the size of the array, the impedance of the low melting point metal 3D electrode and the wet electrode converges. We also measured the stretching range of the low melting point electrode to ensure its stretchability in bio-electrical signal monitoring. We used the 4 × 4 low melting point electrode array with the PDMS substrate. The height of the electrode was 285 μm and the width was 99 μm. We stretched the PDMS substrate of the electrode by the edge in single axis and measured its resistance afterwards. The initial length of the substrate was 1.4 cm (Fig. 3E). It was not conductive when stretched to more than 2 cm, which indicated that the electrode can be stretched by a maximum of 42%. In addition, multi-channel microelectrode arrays can also be developed by applying this method. First, the liquid metal conductor (Ga/ In alloy) was printed by a liquid metal printer on the PVC substrate, then PDMS pre-polymer was coated on the circuit. After being cured at 80 °C in the oven for 1 h, the circuit on the PVC was transferred to the PDMS substrate (Fig. 4A). Afterwards, we aligned the syringe with the connection point in turn and dispensed the microelectrode array (Fig. 4B). Finally, FFC was placed on the connector pad and the PDMS pre-polymer was coated on the liquid metal cured at room temperature for 48 h (Fig. 4C–D). By this approach, the multi-channel flexible electrode was successfully fabricated (Fig. 4E).
were controlled by a computer. Since the alloy is in solid state at room temperature, we fitted an aluminum alloy cylinder heated by constantan wire of resistance 62 Ώ/m outside the plastic syringe to melt the alloy. By supplying electrical power to the constantan resistance wire, the metal was heated from 75 °C to 90 °C, which was monitored by a data collecting instrument (Agilent 34972A, Agilent Technologies Inc., USA). A nitrogen cylinder was used to supply a constant pressure, which was regulated by a solenoid valve on the syringe. In addition, a paint gun and air bottle were used to spray the Ga/In alloy on the surface of the electrode array. Impedance experiment: The impedance data was collected by an electrochemical workstation (Princeton PARSTAT 4000, USA) to assess the electrical properties of the microelectrode. We measured the impedance of the microelectrode and the wet gel electrode. Two shielded copper wires linked with Ga/In alloy conductive film by a clamp, were used to provide reliable connection between the electrode and the equipment. Two microelectrodes were placed on the subject's forearm with 5 cm between them. 10 mV root mean square (RMS) sinusoidal voltage with a frequency range from 1000 Hz to 10 Hz was provided for impedance measurement. Impedance of the Ag/AgCl wet electrode with 0.8 cm diameter was also measured in the same conditions, for comparison. 3. Results and discussion A four-element alloy Bi35In48.6Sn15.9Zn0.4 (melting point: 57.5 °C) was used as the raw material for the electrode. The melting enthalpy and the specific heat capacity of Bi35In48.6Sn15.9Zn0.4 are 28.94 J/g and 0.262 J/(g*°C), respectively, which are much lower than those of commonly used metals. These properties enable its easy liquid-solid phase transition during the dispensing process. After being heated to liquid state, the alloy was pumped through the syringe needle, which was then lifted after the liquid alloy shared adequate contact area with the flexible substrate (Fig. 1A–B). At room temperature, the low melting point metal alloy cools down immediately in to a droplet shape with sharp needle (Fig. 1C). By applying this approach, a low melting point metal can be used to rapidly fabricate flexible 3D microelectrode arrays. The process begins with the PDMS being mixed with a curing agent in the ratio of 10:1, after which the mixture was degassed in a vacuum chamber for 30 min. The pre-polymer was heated at 80 °C for 1 h to solidify and form the substrate of the electrode. Then, the dispensing system was prepared to an adequate temperature to melt the alloy. By setting parameters in the computer program controlling the dispensing process, the droplet-shaped metal with sharp needle can be dispensed to form any size array (Fig. 1D). To make the separated electrodes conductive, Ga/In alloy in liquid state at room temperature was sprayed on the surface of the electrode array (Fig. 1E). Finally, a flexible flat cable (FFC) was placed on the surface of the conductive layer and PDMS prepolymer coated on it followed by curing at room temperature for 48 h, to achieve electrical isolation (Fig. 1F–G). The 4 × 4 microelectrode array before and after sealing was observed through a camera lens and CCD (IMAVISION MER-U3x, China), respectively (Fig. 2A–B). The PDMS sealing layer fixed the electrode on the PDMS substrate and prevented direct contact between the Ga/In alloy and skin. The 6 × 6 microelectrode array was successfully fabricated (Fig. 2C) within 5 min, which suggests fast fabrication and easy shape control of the array using our approach. Generally, the needles of the 3D microelectrode should pierce the stratum corneum (40–100 μm) to lower the skin-electrode impedance but avoid reaching the epidermal layer (200–400 μm), which would cause pain, thus, the height of the 3D microelectrode should be in the range of 100–400 μm. The specific dimension of the electrode was measured by scanning electron microscope (Fig. 2D–F). The average length of the needles was found to be 269.985 ± 25.569 μm, which satisfies the transdermal requirement.
4. Conclusions In summary, we succeeded in directly fabricating a 3D microelectrode using phase transition method. This technology enables the rapid formation of a flexible 3D microelectrode array because of its transition mechanics. The dimension of the electrode and the size of the array can be easily controlled by just setting the system parameters including lifting velocity of the syringe needle and sample temperature. By increasing the sample temperature, the height of the electrode can reach up to 1 mm. It may also be noted that the height of the sharp needle of the electrode can be adjusted to penetrate just up to the stratum corneum, which decreases the skin-electrode impedance while protecting 737
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Fig. 2. The microelectrode array as observed through a camera lens and CCD. (A) The unsealed 4 × 4 microelectrode array. (B) The 4 × 4 3D microelectrode array with PDMS insulation layer. (C) A 6 × 6 microelectrode array as observed through a scanning electron microscope (SEM). (D) The microelectrode array. (E) The single needle of the electrode. (F) The pin point of the electrode.
Fig. 3. (A) The relationship between the height of the electrode and lifting velocity at different sample temperatures. (B) The relationship between the width of the electrode and lifting velocity at different sample temperatures. (C) The relationship between the depth-width ratio of the electrode and lifting velocity at different sample temperatures. (D) The skin-electrode impedance of the low melting metal electrode and Ag/AgCl electrode at frequency ranging from 10 Hz to 1000 Hz, respectively. (E) The relationship between the resistance of the low melting point metal electrode and the length of the electrode pad.
practical issues to be solved in the near future.
the patient from feeling any pain. Apart from applying in acquiring bioelectric signals, the 3D microelectrode has strong potential in bio-sensing and preparation of electrode templates because of its flexibility and fast fabrication characteristic. Our subsequent research will aim to acquire ECG and EEG signals using our electrodes. As a completely new fabrication method, our research raises important fundamental and
Acknowledgements This research was supported by the National Natural Science Foundation of China under Grant No. 81471749, the National Key R&D 738
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Fig. 4. (A)–(D) Schematic illustration of the fabrication process of multi-channel flexible microelectrode array device. (E) Image of multi-channel flexible microelectrode.
Program of China under Grant No. 2017YFC0803507, the Tsinghua University Initiative Scientific Research Program under Grant No. 2014Z01001, and the Foundation of Beijing Laboratory in Biomedical Technology and Instruments. We express gratitude to Capital Bio Corporation Micro Systems Engineering Lab for their help.
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