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Rigid and Flexible Organic Electrochemical Transistor Arrays for Monitoring Action Potentials from Electrogenic Cells Chunlei Yao, Qianqian Li, Jing Guo, Feng Yan, and I-Ming Hsing* The advancement of neuroscience and cardiology has greatly improved our understanding of neurodegenerative and cardiovascular diseases. Yet it also poses challenges on the development of bioelectronics interfacing with organ-specific cells/ tissues.[1] Specifically, a new device platform would require not only more accurate measurements of electrophysiological characteristics (e.g., action potential) but also new modalities, such as elasticity and versatility to meet emerging needs. Much effort has been devoted to applying new materials or novel architectures onto conventional techniques. For example, microelectrode array (MEA), a gold standard technique in this area, has evolved from bare planar surface to various surface coatings[2,3] and nanostructures[4,5] in order to reduce impedance between electrolyte and electrode surface. A conformable MEA fabricated on flexible parylene substrate for in vivo measurements is another recent example illustrating this direction.[6] Notably, a variety of semiconductors and device architectures, including Si nanowire,[7–9] diamond,[10,11] graphene[12–14] and 3D field-effect transistors (FETs),[15–17] have also been intensively studied to improve the performance of transistor-based devices since the first report of silicon-based transistor for action potential measurement in 1991.[18] Organic electronic materials (e.g., conducting polymers)based transistors, particularly organic electrochemical transistors (OECTs), have recently received lots of attentions in this trend.[19] OECTs have found niches in biological applications with several attractive characteristics.[20–22] Unlike FET-based devices requiring dielectric layers for electrostatic gating, OECTs have the organic active layers directly interfaced with cells or other biological systems in electrolytes,[23] whose channel conductivities are electrochemically modulated through ion doping by extracellular matrix/electrolytes. The
C. Yao, Dr. Q. Li, Dr. J. Guo, Prof. I.-M. Hsing Bioengineering Graduate Program Division of Biomedical Engineering The Hong Kong University of Science and Technology Hong Kong, China E-mail: [email protected] Prof. F. Yan Department of Applied Physics and Materials Research Centre The Hong Kong Polytechnic University Hong Kong, China Prof. I.-M. Hsing Department of Chemical and Biomolecular Engineering The Hong Kong University of Science and Technology Hong Kong, China
DOI: 10.1002/adhm.201400406
Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
combination of dielectric-free interface and electrochemical operation of OECTs results in that the whole thickness of the porous conducting polymer layers, instead of only the dielectric surfaces in FETs, are involved in electrochemical interactions with ions, which makes OECTs ideal devices for coupling biological ion currents (frequency normally below 10 kHz) to electronic ones. Our recent work has demonstrated such coupling by using OECT for recording transepithelial ion transport, in which a physiological signal was readily converted to an electrical readout.[24] Properties of the OECT's organic material surface, which is in direct contact with cells and other biological systems, could also be easily tailored to facilitate various biological applications without disrupting its electronic performance.[25,26] Moreover, from the standpoint of device fabrication and material adaptability, organic electronic materials are compatible for fabrication onto flexible substrates, the feature of which is critical for in vivo study of tissues or organs to provide mechanical compliance. A very recent work by Malliaras and co-workers[27] has vividly shown the merits of a flexible OECT device for brain activity research. The compatibility on flexible substrates is not only useful for in vivo applications. In fact, flexible devices could conformably adapt to nonplanar forms to monitor cells in 3D microenvironment.[28] Furthermore, it is anticipated that OECT arrays fabricated on plastics by printing technology, like roll-to-roll printing, would largely reduce the manufacturing cost of the in vitro monitoring chips currently dominated by glass-based MEA devices.[29] Despite these intriguing attributes provided by OECTs, the direct application of OECTs to interface with excitable cells, to our knowledge, has not been realized. Therefore, a study of OECT's performance for in vitro usage, in particular monitoring action potentials, would benefit the development of bioelectronics by taking full advantage of organic electronic materials. Here, we present the first study of using OECT arrays for monitoring action potentials from cardiomyocyte (i.e., cardiac muscle cell)-like HL-1 cells on both rigid and flexible substrates with improved signal qualities compared to those recorded from grapheme-based electrolyte-gated FETs.[13] An excellent signal to noise ratio (S/N) was achieved while signal amplitudes could be tuned by varying gate voltages. The feasibility of using OECT arrays for drug evaluation was also demonstrated in this work. Moreover, the results on flexible devices showed that negligible changes in the performance of flexible OECT arrays could be observed when they were bent to either side. And the action potential derived signals could be recorded with similar signal quality to that from devices on rigid substrates. Figure 1a shows the schematic of using an OECT for action potential monitoring. A cleft is formed between cells and the
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Figure 1. a) Schematic view of HL-1 cells integrated with an OECT. b) Optical image showing eight OECTs in the center of an OECT array. The active area of an individual OECT is 40 × 30 µm2. Scale bar is 100 µm. c) Transfer characteristics of four OECTs in the same array measured in cell culture medium with confluent HL-1 cells layer on. Vds = −0.4 V. d) Corresponding transconductance curves derived from transfer characteristics in (c). Red dash line indicates the working point of OECTs for signal acquisition.
OECT. By implementing an extracellular detection scheme, local field potential changes caused by ion flows of action potentials are detected by the OECT beneath through modulating its channel current. We firstly applied OECTs for monitoring action potentials on rigid substrates. An OECT array with its central part shown in Figure 1b was fabricated on a glass substrate. Each OECT had an active area of 30 × 40 µm2 and was separated by a distance of 200 µm from each other. The fabrication process began with defining the gold source/drain electrodes and interconnects by photolithography, followed by a SU-8 layer deposition for surface passivation. Biocompatible conducting polymer, poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonic acid) (PEDOT:PSS) was then deposited between the source and drain to form a functional OECT device. Details of the OECT array configuration and fabrication process could be found in the Experimental Section (Supporting Information). The as-fabricated OECT array was then characterized in cell culture medium solution with Ag/AgCl wire as gate electrode. Figure 1c shows the transfer characteristics of four OECTs in the same chip for drain voltage Vds = −0.4 V and gate voltage Vg swept from −0.4 to 0.4 V. These characteristics demonstrated a typical low voltage (below 1 V) operation in the depletion mode, which is the characteristic of PEDOT:PSS-based OECT. The corresponding transconductance curves derived from those transfer characteristics are shown in Figure 1d. It is notable that the high transconductances up to about 2.5 mS could be achieved. This result is consistent with a previous study[30] and favorable for the transduction of action potential induced signals, because the higher local in situ amplification would lead
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to the better quality of output signals. Moreover, despite some minor deviations, the highest transconductance for all the measurements was obtained at a gate voltage around −0.2 V. An earlier study has shown that the gate voltage, at which an OECT achieved maximal transconductance, could be tuned by varying film thickness and the channel width to length ratio.[31] In our study, such gate voltage was chosen as the working gate voltage (−0.2 V) for monitoring action potential to achieve maximal output. We used an established cardiomycyte-like cell line HL-1 as the study model here. Atrial cardiac muscle cell-derived HL-1 is well known for its capacity to maintain the cardiac-specific phenotype and spontaneous contraction during the culture period,[32] which provides a good platform to evaluate the OECTs for action potential measurements. To do so, the HL-1 cells were seeded into the glass ring culture chamber glued on the top of OECT array surface, as shown in Figure 2a. After 3 to 4 days’ culturing, HL-1 cells formed a dense confluent cell layer covering the surface of OECTs, as it can be seen in Figure 2b. Individual or cluster of cells started to contract spontaneously after confluency reached, to a large extent. Figure 2c shows the traces of regularly spaced spikes recorded on eight OECTs on the same chip. To eliminate the impact of transconductance variances among different OECTs (see Figure 1d) on signal recording, recorded current signals were converted to equivalent gate voltage changes by using transconductance curves for a uniform comparison. Using a home-built amplification and recording circuit (Supporting Information), a noise level (peak to peak) in the range of 100–200 µV could be achieved (Supporting Information), which was low enough
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COMMUNICATION Figure 2. a) Photograph of the whole device, the size is similar to a coin of 5 HK$. b) Microscopy image (phase contrast) of confluent HL-1 cell layer grew on an OECT array. c) Representative traces of spikes recorded by eight OECTs in the same chip. Current traces were converted to voltage traces by using transconductance curves for uniform comparison. Varied amplitudes of signals were recorded by OECTs at different sites with no signal detected by OECT1. The heterogeneity of signal amplitudes was caused by different cleft coverage and seal resistances between cells and OECTs. d) Shapes of single spikes corresponded with signals showing in (c). The heterogeneity of signal shapes was due to different expression levels of ion channels on the cleft side’s cell membrane.
to differentiate cardiac action potential signals. Even with no optimization on the recording system, our OECT devices with high transconductances acquired results with excellent quality, supported by the recorded spikes with S/N values routinely larger than 4 (see Figure 2 for examples) and even larger than 10 in some scenarios (Supporting information). Such high S/N results are comparable to those reported previously using silicon nanowire transistors[9] and that from organic cell stimulating and sensing transistor device[33] for primary neurons’ sensing and stimulation. It should be noted that much higher S/N values were demonstrated formerly by nanocavity electrode arrays[34] and nanopillar electrodes.[35] However, compared with the straightforward fabrication processes of OECT, sophisticated nanofabrication procedures, like focused ion beam etching, were heavily involved in those nanodevices’ manufacturing, which hindered them from being widely implemented. Figure 2d depicts the corresponding individual spikes enlarged from Figure 2c. Varied shapes and amplitudes can be found among different OECTs. The heterogeneity in signals can be explained by the so-called point-contact model described previously.[36] Briefly, signal shape is determined by the different expression levels of several ion channels on the cleft side’s cell membrane, while the signal amplitude is mainly decided by the cleft coverage and seal resistance. One important feature of the OECT devices needs to be pointed out here is that the signal amplitude from monitoring action potential can be tuned by varying the gate voltage. For OECT, extracellular potential change induced by action potential activities was recorded and converted to channel current change with the converting capability calibrated by transconductance. By varying gate voltages (i.e., changing transconductances),
Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
amplitudes of recorded current change could be modulated. Figure 3a shows the three recorded current traces of spikes from one OECT in the same experiment at different gate voltages. Modulation of peak to peak amplitude with a factor of about 3 was achieved for current amplitudes of 1.83 ± 0.04, 1.45 ± 0.04, and 0.63 ± 0.02 µA recorded at Vg = −0.2, 0 and +0.2 V, respectively. Notably, calibrated peak to peak voltage changes by using transconductance curves remained relatively constant, as shown in Figure 3b. These results demonstrate that the in situ amplification capacity of OECT is controlled by gate voltage, and also reconfirm the necessity to convert the recorded current signals to equivalent voltage signals for accurate comparisons among data from different OECT devices. The feasibility of using OECT arrays for chronotropic (i.e., changing heart rate) drug screening was also conducted in this work. A well-established positive chronotropic agent, isoproterenol (see Supporting Information for mechanism), was used to stimulate the cells in basic culture medium without any other supplement (see Experimental Section, Supporting Information). Figure 3c shows the representative traces of spikes recorded before and after the treatment of 1 × 10−3 M isoproterenol with the frequency increased of 116.28 ± 18.52%. Traces of spikes for other three concentrations can be found in Supporting Information. Figure 3d summarizes the data for action potential frequency changes due to the different isoproterenol concentrations. For 0.1 × 10−3 and 0.01 × 10−3 M of isoproterenol, action potential frequency recorded by OECTs increased 92.2 ± 5.81% and 52.29 ± 5.41%, respectively. However, the frequency decreased by 27.83 ± 6.27% when isoproterenol’s concentration was lowered down to 0.001 × 10−3 M. This result can be explained by that such low concentration of isoproterenol is
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Figure 3. a) Recorded current traces at gate voltage of +0.2, 0, and −0.2 V, respectively. Corresponding transconductances are 2.57, 2.03, and 0.95 mS, respectively. b) Summary table of gate voltage’s effect on recorded signals. Open circles are the recorded peak to peak current changes. Open squares are corresponding to the calibrated voltage changes. Data were obtained from the same experiment in (a). c) Representative recorded traces of spikes before and after treatment of 1 × 10−3 M isoproterenol. The frequency increased 116.28 ± 18.52% for 1 × 10−3 M isoproterenol. d) Summary of the recorded frequency changes after treated with 1 × 10−3, 0.1 × 10−3, 0.01 × 10−3, and 0.001 × 10−3 M isoproterenol, respectively. The larger concentration of isoproterenol used, the higher frequency from the HL-1 beating behavior can be detected.
not sufficient to maintain the beating frequency induced by the 0.1 × 10−3 M of hormone named norepinephrine prior existing in the HL-1 routine culture (see Supporting Information), which is essential to maintain cell spontaneous contraction. Nevertheless, at the same concentration level, the effect of isoproterenol surpasses that of norepinephrine, which is found to be consistent with the previous studies.[37] This quantitative analysis of drug efficacy on cell behavior was facilitated by our OECT device, which suggests its promising applications in evaluating the excitable cell-targeted drug and investigating the unintended effects on heart by chemicals. OECTs for monitoring action potentials were then performed on flexible substrates. We fabricated flexible OECT arrays on polyethylene terephthalate (PET) substrates through similar procedures of that on glass substrates. To keep the whole device flexible, cell culture chamber made of poly (dimethylsiloxane) (PDMS), instead of glass ring, was glued onto the surface by silicone. The as-fabricated device could be bent to both sides with no visible damages caused, as shown in Figure 4a. Moreover, the transfer characteristics in Figure 4b show that a flexible OECT could maintain a stable performance under different bending strains provided by the flexible nature of PEDOT:PSS, which is similar to that of an flexible OECT-based DNA sensor previously reported by our group.[38] HL-1 cells were seeded onto the surface of the flexible OECT array and after 3–4 days’
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culture, a healthy confluent layer was formed (Supporting Information) with clusters of cells showing spontaneous contraction. Figure 4c depicts the traces of spikes induced by action potential and Figure 4d shows the enlarged single spikes recorded on a flexible OECT device. Similar noise level to that on a rigid OECT device has been achieved in flexible one with a maximal noise of around 200 µV, as it can be seen in the upper trace of Figure 4d. Excellent signal to noise ratio was maintained on the flexible OECT array with values routinely larger than 4. The success of using flexible OECT arrays on plastics to monitor action potential paves the way for realizing cheap, disposable in vitro monitoring chips. With its stable performance under strains, it also opens up the possibility to apply mechanical cues on cells with simultaneous electrophysiological sensing, which would be useful for mechanobiology study.[39,40] To sum up, for the first time, cardiac action potential was successfully monitored by OECTs on both rigid and flexible substrates. Due to high transconductances, excellent signal to noise ratios (routinely larger than 4) were achieved. The feasibility of using OECT arrays for drug screening was also demonstrated by quantitatively recording the positive chronotropic effect of isoproterenol. Thanks to its sensitive signal acquisition ability and unique feature on flexibility, OECTs show a good potential to be a novel in vitro monitoring platform serving both conventional tasks and emerging needs.
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Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
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COMMUNICATION Figure 4. a) Photographs of a flexible OECT array device bent to both sides. b) Transfer characteristics of a flexible OECT measured at different bending status in cell culture medium. Negligible change was observed among different bending status. Representative traces of spikes (c) and corresponding zoomed single spikes (d) recorded on the flexible OECT array.
Supporting Information Supporting Information is available from the Wiley Online Library or from the author.
Acknowledgements C.Y. and Q.L. contributed equally to this work. The authors thank financial support from the Theme-based Research Scheme from Research Grant Council (RGC) of the Hong Kong SAR Government, China (Project number: T13–706/11–2). Received: July 14, 2014 Revised: September 25, 2014 Published online:
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[4] D. Bruggemann, B. Wolfrum, V. Maybeck, Y. Mourzina, M. Jansen, A. Offenhausser, Nanotechnology 2011, 22, 265104. [5] J. H. Kim, G. Kang, Y. Nam, Y. K. Choi, Nanotechnology 2010, 21, 85303. [6] D. Khodagholy, T. Doublet, M. Gurfinkel, P. Quilichini, E. Ismailova, P. Leleux, T. Herve, S. Sanaur, C. Bernard, G. G. Malliaras, Adv. Mater. 2011, 23, H268. [7] F. Patolsky, B. P. Timko, G. Yu, Y. Fang, A. B. Greytak, G. Zheng, C. M. Lieber, Science 2006, 313, 1100. [8] J. F. Eschermann, R. Stockmann, M. Hueske, X. T. Vu, S. Ingebrandt, A. Offenhäusser, Appl. Phys. Lett. 2009, 95, 083703. [9] T. Cohen-Karni, B. P. Timko, L. E. Weiss, C. M. Lieber, Proc. Natl. Acad. Sci. U.S.A. 2009, 106, 7309. [10] M. Dankerl, S. Eick, B. Hofmann, M. Hauf, S. Ingebrandt, A. Offenhäusser, M. Stutzmann, J. A. Garrido, Adv. Funct. Mater. 2009, 19, 2915. [11] V. Maybeck, R. Edgington, A. Bongrain, J. O. Welch, E. Scorsone, P. Bergonzo, R. B. Jackman, A. Offenhausser, Adv. Healthcare Mater. 2014, 3, 283. [12] T. Cohen-Karni, Q. Qing, Q. Li, Y. Fang, C. M. Lieber, Nano Lett. 2010, 10, 1098. [13] L. H. Hess, M. Jansen, V. Maybeck, M. V. Hauf, M. Seifert, M. Stutzmann, I. D. Sharp, A. Offenhausser, J. A. Garrido, Adv. Mater. 2011, 23, 5045.
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www.MaterialsViews.com [14] Z. Cheng, J. Hou, Q. Zhou, T. Li, H. Li, L. Yang, K. Jiang, C. Wang, Y. Li, Y. Fang, Nano Lett. 2013, 13, 2902. [15] B. Tian, T. Cohen-Karni, Q. Qing, X. Duan, P. Xie, C. M. Lieber, Science 2010, 329, 830. [16] X. Duan, R. Gao, P. Xie, T. Cohen-Karni, Q. Qing, H. S. Choe, B. Tian, X. Jiang, C. M. Lieber, Nat. Nanotechnol. 2012, 7, 174. [17] T. M. Fu, X. Duan, Z. Jiang, X. Dai, P. Xie, Z. Cheng, C. M. Lieber, Proc. Natl. Acad. Sci. U.S.A. 2014, 111, 1259. [18] P. Fromherz, A. Offenhausser, T. Vetter, J. Weis, Science 1991, 252, 1290. [19] M. Berggren, A. Richter-Dahlfors, Adv. Mater. 2007, 19, 3201. [20] R. M. Owens, G. G. Malliaras, MRS Bull. 2010, 35, 449. [21] G. G. Malliaras, Biochim. Biophys. Acta 2013, 1830, 4286. [22] P. Lin, F. Yan, Adv. Mater. 2012, 24, 34. [23] P. Lin, F. Yan, J. Yu, H. L. Chan, M. Yang, Adv. Mater. 2010, 22, 3655. [24] C. Yao, C. Xie, P. Lin, F. Yan, P. Huang, I. M. Hsing, Adv. Mater. 2013, 25, 6575. [25] R.-X. He, M. Zhang, F. Tan, P. H. M. Leung, X.-Z. Zhao, H. L. W. Chan, M. Yang, F. Yan, J. Mater. Chem. 2012, 22, 22072. [26] X. Strakosas, M. Sessolo, A. Hama, J. Rivnay, E. Stavrinidou, G. G. Malliaras, R. M. Owens, J. Mater. Chem. B 2014, 22537. [27] D. Khodagholy, T. Doublet, P. Quilichini, M. Gurfinkel, P. Leleux, A. Ghestem, E. Ismailova, T. Herve, S. Sanaur, C. Bernard, G. G. Malliaras, Nat. Commun. 2013, 4, 1575.
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[28] B. Tian, J. Liu, T. Dvir, L. Jin, J. H. Tsui, Q. Qing, Z. Suo, R. Langer, D. S. Kohane, C. M. Lieber, Nat. Mater. 2012, 11, 986. [29] J. Rivnay, R. M. Owens, G. G. Malliaras, Chem. Mater. 2014, 26, 679. [30] D. Khodagholy, J. Rivnay, M. Sessolo, M. Gurfinkel, P. Leleux, L. H. Jimison, E. Stavrinidou, T. Herve, S. Sanaur, R. M. Owens, G. G. Malliaras, Nat. Commun. 2013, 4, 2133. [31] J. Rivnay, P. Leleux, M. Sessolo, D. Khodagholy, T. Herve, M. Fiocchi, G. G. Malliaras, Adv. Mater. 2013, 25, 7010. [32] W. C. Claycomb, N. A. Lanson, B. S. Stallworth, D. B. Egeland, J. B. Delcarpio, A. Bahinski, N. J. Izzo, Proc. Natl. Acad. Sci. U.S.A. 1998, 95, 2979. [33] V. Benfenati, S. Toffanin, S. Bonetti, G. Turatti, A. Pistone, M. Chiappalone, A. Sagnella, A. Stefani, G. Generali, G. Ruani, D. Saguatti, R. Zamboni, M. Muccini, Nat. Mater. 2013, 12, 672. [34] B. Hofmann, E. Katelhon, M. Schottdorf, A. Offenhausser, B. Wolfrum, Lab Chip 2011, 11, 1054. [35] C. Xie, Z. Lin, L. Hanson, Y. Cui, B. Cui, Nat. Nanotechnol. 2012, 7, 185. [36] P. Fromherz, Eur. Biophys. J. 1999, 28, 254. [37] C.-K. Yeung, S. Ingebrandt, M. Krause, A. Offenhäusser, W. Knoll, J. Pharmacol. Toxicol. Meth. 2001, 45, 207. [38] P. Lin, X. Luo, I. M. Hsing, F. Yan, Adv. Mater. 2011, 23, 4035. [39] W. J. Tyler, Nat. Rev. Neurosci. 2012, 13, 867. [40] K. Takahashi, Y. Kakimoto, K. Toda, K. Naruse, J. Cell. Mol. Med. 2013, 17, 225.
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Rigid and Flexible Organic Electrochemical Transistor Arrays for Monitoring Action Potentials from Electrogenic Cells Chunlei Yao, Qianqian Li, Jing Guo, Feng Yan, and I-Ming Hsing* The advancement of neuroscience and cardiology has greatly improved our understanding of neurodegenerative and cardiovascular diseases. Yet it also poses challenges on the development of bioelectronics interfacing with organ-specific cells/ tissues.[1] Specifically, a new device platform would require not only more accurate measurements of electrophysiological characteristics (e.g., action potential) but also new modalities, such as elasticity and versatility to meet emerging needs. Much effort has been devoted to applying new materials or novel architectures onto conventional techniques. For example, microelectrode array (MEA), a gold standard technique in this area, has evolved from bare planar surface to various surface coatings[2,3] and nanostructures[4,5] in order to reduce impedance between electrolyte and electrode surface. A conformable MEA fabricated on flexible parylene substrate for in vivo measurements is another recent example illustrating this direction.[6] Notably, a variety of semiconductors and device architectures, including Si nanowire,[7–9] diamond,[10,11] graphene[12–14] and 3D field-effect transistors (FETs),[15–17] have also been intensively studied to improve the performance of transistor-based devices since the first report of silicon-based transistor for action potential measurement in 1991.[18] Organic electronic materials (e.g., conducting polymers)based transistors, particularly organic electrochemical transistors (OECTs), have recently received lots of attentions in this trend.[19] OECTs have found niches in biological applications with several attractive characteristics.[20–22] Unlike FET-based devices requiring dielectric layers for electrostatic gating, OECTs have the organic active layers directly interfaced with cells or other biological systems in electrolytes,[23] whose channel conductivities are electrochemically modulated through ion doping by extracellular matrix/electrolytes. The
C. Yao, Dr. Q. Li, Dr. J. Guo, Prof. I.-M. Hsing Bioengineering Graduate Program Division of Biomedical Engineering The Hong Kong University of Science and Technology Hong Kong, China E-mail: [email protected] Prof. F. Yan Department of Applied Physics and Materials Research Centre The Hong Kong Polytechnic University Hong Kong, China Prof. I.-M. Hsing Department of Chemical and Biomolecular Engineering The Hong Kong University of Science and Technology Hong Kong, China
DOI: 10.1002/adhm.201400406
Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
combination of dielectric-free interface and electrochemical operation of OECTs results in that the whole thickness of the porous conducting polymer layers, instead of only the dielectric surfaces in FETs, are involved in electrochemical interactions with ions, which makes OECTs ideal devices for coupling biological ion currents (frequency normally below 10 kHz) to electronic ones. Our recent work has demonstrated such coupling by using OECT for recording transepithelial ion transport, in which a physiological signal was readily converted to an electrical readout.[24] Properties of the OECT's organic material surface, which is in direct contact with cells and other biological systems, could also be easily tailored to facilitate various biological applications without disrupting its electronic performance.[25,26] Moreover, from the standpoint of device fabrication and material adaptability, organic electronic materials are compatible for fabrication onto flexible substrates, the feature of which is critical for in vivo study of tissues or organs to provide mechanical compliance. A very recent work by Malliaras and co-workers[27] has vividly shown the merits of a flexible OECT device for brain activity research. The compatibility on flexible substrates is not only useful for in vivo applications. In fact, flexible devices could conformably adapt to nonplanar forms to monitor cells in 3D microenvironment.[28] Furthermore, it is anticipated that OECT arrays fabricated on plastics by printing technology, like roll-to-roll printing, would largely reduce the manufacturing cost of the in vitro monitoring chips currently dominated by glass-based MEA devices.[29] Despite these intriguing attributes provided by OECTs, the direct application of OECTs to interface with excitable cells, to our knowledge, has not been realized. Therefore, a study of OECT's performance for in vitro usage, in particular monitoring action potentials, would benefit the development of bioelectronics by taking full advantage of organic electronic materials. Here, we present the first study of using OECT arrays for monitoring action potentials from cardiomyocyte (i.e., cardiac muscle cell)-like HL-1 cells on both rigid and flexible substrates with improved signal qualities compared to those recorded from grapheme-based electrolyte-gated FETs.[13] An excellent signal to noise ratio (S/N) was achieved while signal amplitudes could be tuned by varying gate voltages. The feasibility of using OECT arrays for drug evaluation was also demonstrated in this work. Moreover, the results on flexible devices showed that negligible changes in the performance of flexible OECT arrays could be observed when they were bent to either side. And the action potential derived signals could be recorded with similar signal quality to that from devices on rigid substrates. Figure 1a shows the schematic of using an OECT for action potential monitoring. A cleft is formed between cells and the
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Figure 1. a) Schematic view of HL-1 cells integrated with an OECT. b) Optical image showing eight OECTs in the center of an OECT array. The active area of an individual OECT is 40 × 30 µm2. Scale bar is 100 µm. c) Transfer characteristics of four OECTs in the same array measured in cell culture medium with confluent HL-1 cells layer on. Vds = −0.4 V. d) Corresponding transconductance curves derived from transfer characteristics in (c). Red dash line indicates the working point of OECTs for signal acquisition.
OECT. By implementing an extracellular detection scheme, local field potential changes caused by ion flows of action potentials are detected by the OECT beneath through modulating its channel current. We firstly applied OECTs for monitoring action potentials on rigid substrates. An OECT array with its central part shown in Figure 1b was fabricated on a glass substrate. Each OECT had an active area of 30 × 40 µm2 and was separated by a distance of 200 µm from each other. The fabrication process began with defining the gold source/drain electrodes and interconnects by photolithography, followed by a SU-8 layer deposition for surface passivation. Biocompatible conducting polymer, poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulfonic acid) (PEDOT:PSS) was then deposited between the source and drain to form a functional OECT device. Details of the OECT array configuration and fabrication process could be found in the Experimental Section (Supporting Information). The as-fabricated OECT array was then characterized in cell culture medium solution with Ag/AgCl wire as gate electrode. Figure 1c shows the transfer characteristics of four OECTs in the same chip for drain voltage Vds = −0.4 V and gate voltage Vg swept from −0.4 to 0.4 V. These characteristics demonstrated a typical low voltage (below 1 V) operation in the depletion mode, which is the characteristic of PEDOT:PSS-based OECT. The corresponding transconductance curves derived from those transfer characteristics are shown in Figure 1d. It is notable that the high transconductances up to about 2.5 mS could be achieved. This result is consistent with a previous study[30] and favorable for the transduction of action potential induced signals, because the higher local in situ amplification would lead
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to the better quality of output signals. Moreover, despite some minor deviations, the highest transconductance for all the measurements was obtained at a gate voltage around −0.2 V. An earlier study has shown that the gate voltage, at which an OECT achieved maximal transconductance, could be tuned by varying film thickness and the channel width to length ratio.[31] In our study, such gate voltage was chosen as the working gate voltage (−0.2 V) for monitoring action potential to achieve maximal output. We used an established cardiomycyte-like cell line HL-1 as the study model here. Atrial cardiac muscle cell-derived HL-1 is well known for its capacity to maintain the cardiac-specific phenotype and spontaneous contraction during the culture period,[32] which provides a good platform to evaluate the OECTs for action potential measurements. To do so, the HL-1 cells were seeded into the glass ring culture chamber glued on the top of OECT array surface, as shown in Figure 2a. After 3 to 4 days’ culturing, HL-1 cells formed a dense confluent cell layer covering the surface of OECTs, as it can be seen in Figure 2b. Individual or cluster of cells started to contract spontaneously after confluency reached, to a large extent. Figure 2c shows the traces of regularly spaced spikes recorded on eight OECTs on the same chip. To eliminate the impact of transconductance variances among different OECTs (see Figure 1d) on signal recording, recorded current signals were converted to equivalent gate voltage changes by using transconductance curves for a uniform comparison. Using a home-built amplification and recording circuit (Supporting Information), a noise level (peak to peak) in the range of 100–200 µV could be achieved (Supporting Information), which was low enough
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COMMUNICATION Figure 2. a) Photograph of the whole device, the size is similar to a coin of 5 HK$. b) Microscopy image (phase contrast) of confluent HL-1 cell layer grew on an OECT array. c) Representative traces of spikes recorded by eight OECTs in the same chip. Current traces were converted to voltage traces by using transconductance curves for uniform comparison. Varied amplitudes of signals were recorded by OECTs at different sites with no signal detected by OECT1. The heterogeneity of signal amplitudes was caused by different cleft coverage and seal resistances between cells and OECTs. d) Shapes of single spikes corresponded with signals showing in (c). The heterogeneity of signal shapes was due to different expression levels of ion channels on the cleft side’s cell membrane.
to differentiate cardiac action potential signals. Even with no optimization on the recording system, our OECT devices with high transconductances acquired results with excellent quality, supported by the recorded spikes with S/N values routinely larger than 4 (see Figure 2 for examples) and even larger than 10 in some scenarios (Supporting information). Such high S/N results are comparable to those reported previously using silicon nanowire transistors[9] and that from organic cell stimulating and sensing transistor device[33] for primary neurons’ sensing and stimulation. It should be noted that much higher S/N values were demonstrated formerly by nanocavity electrode arrays[34] and nanopillar electrodes.[35] However, compared with the straightforward fabrication processes of OECT, sophisticated nanofabrication procedures, like focused ion beam etching, were heavily involved in those nanodevices’ manufacturing, which hindered them from being widely implemented. Figure 2d depicts the corresponding individual spikes enlarged from Figure 2c. Varied shapes and amplitudes can be found among different OECTs. The heterogeneity in signals can be explained by the so-called point-contact model described previously.[36] Briefly, signal shape is determined by the different expression levels of several ion channels on the cleft side’s cell membrane, while the signal amplitude is mainly decided by the cleft coverage and seal resistance. One important feature of the OECT devices needs to be pointed out here is that the signal amplitude from monitoring action potential can be tuned by varying the gate voltage. For OECT, extracellular potential change induced by action potential activities was recorded and converted to channel current change with the converting capability calibrated by transconductance. By varying gate voltages (i.e., changing transconductances),
Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
amplitudes of recorded current change could be modulated. Figure 3a shows the three recorded current traces of spikes from one OECT in the same experiment at different gate voltages. Modulation of peak to peak amplitude with a factor of about 3 was achieved for current amplitudes of 1.83 ± 0.04, 1.45 ± 0.04, and 0.63 ± 0.02 µA recorded at Vg = −0.2, 0 and +0.2 V, respectively. Notably, calibrated peak to peak voltage changes by using transconductance curves remained relatively constant, as shown in Figure 3b. These results demonstrate that the in situ amplification capacity of OECT is controlled by gate voltage, and also reconfirm the necessity to convert the recorded current signals to equivalent voltage signals for accurate comparisons among data from different OECT devices. The feasibility of using OECT arrays for chronotropic (i.e., changing heart rate) drug screening was also conducted in this work. A well-established positive chronotropic agent, isoproterenol (see Supporting Information for mechanism), was used to stimulate the cells in basic culture medium without any other supplement (see Experimental Section, Supporting Information). Figure 3c shows the representative traces of spikes recorded before and after the treatment of 1 × 10−3 M isoproterenol with the frequency increased of 116.28 ± 18.52%. Traces of spikes for other three concentrations can be found in Supporting Information. Figure 3d summarizes the data for action potential frequency changes due to the different isoproterenol concentrations. For 0.1 × 10−3 and 0.01 × 10−3 M of isoproterenol, action potential frequency recorded by OECTs increased 92.2 ± 5.81% and 52.29 ± 5.41%, respectively. However, the frequency decreased by 27.83 ± 6.27% when isoproterenol’s concentration was lowered down to 0.001 × 10−3 M. This result can be explained by that such low concentration of isoproterenol is
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Figure 3. a) Recorded current traces at gate voltage of +0.2, 0, and −0.2 V, respectively. Corresponding transconductances are 2.57, 2.03, and 0.95 mS, respectively. b) Summary table of gate voltage’s effect on recorded signals. Open circles are the recorded peak to peak current changes. Open squares are corresponding to the calibrated voltage changes. Data were obtained from the same experiment in (a). c) Representative recorded traces of spikes before and after treatment of 1 × 10−3 M isoproterenol. The frequency increased 116.28 ± 18.52% for 1 × 10−3 M isoproterenol. d) Summary of the recorded frequency changes after treated with 1 × 10−3, 0.1 × 10−3, 0.01 × 10−3, and 0.001 × 10−3 M isoproterenol, respectively. The larger concentration of isoproterenol used, the higher frequency from the HL-1 beating behavior can be detected.
not sufficient to maintain the beating frequency induced by the 0.1 × 10−3 M of hormone named norepinephrine prior existing in the HL-1 routine culture (see Supporting Information), which is essential to maintain cell spontaneous contraction. Nevertheless, at the same concentration level, the effect of isoproterenol surpasses that of norepinephrine, which is found to be consistent with the previous studies.[37] This quantitative analysis of drug efficacy on cell behavior was facilitated by our OECT device, which suggests its promising applications in evaluating the excitable cell-targeted drug and investigating the unintended effects on heart by chemicals. OECTs for monitoring action potentials were then performed on flexible substrates. We fabricated flexible OECT arrays on polyethylene terephthalate (PET) substrates through similar procedures of that on glass substrates. To keep the whole device flexible, cell culture chamber made of poly (dimethylsiloxane) (PDMS), instead of glass ring, was glued onto the surface by silicone. The as-fabricated device could be bent to both sides with no visible damages caused, as shown in Figure 4a. Moreover, the transfer characteristics in Figure 4b show that a flexible OECT could maintain a stable performance under different bending strains provided by the flexible nature of PEDOT:PSS, which is similar to that of an flexible OECT-based DNA sensor previously reported by our group.[38] HL-1 cells were seeded onto the surface of the flexible OECT array and after 3–4 days’
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culture, a healthy confluent layer was formed (Supporting Information) with clusters of cells showing spontaneous contraction. Figure 4c depicts the traces of spikes induced by action potential and Figure 4d shows the enlarged single spikes recorded on a flexible OECT device. Similar noise level to that on a rigid OECT device has been achieved in flexible one with a maximal noise of around 200 µV, as it can be seen in the upper trace of Figure 4d. Excellent signal to noise ratio was maintained on the flexible OECT array with values routinely larger than 4. The success of using flexible OECT arrays on plastics to monitor action potential paves the way for realizing cheap, disposable in vitro monitoring chips. With its stable performance under strains, it also opens up the possibility to apply mechanical cues on cells with simultaneous electrophysiological sensing, which would be useful for mechanobiology study.[39,40] To sum up, for the first time, cardiac action potential was successfully monitored by OECTs on both rigid and flexible substrates. Due to high transconductances, excellent signal to noise ratios (routinely larger than 4) were achieved. The feasibility of using OECT arrays for drug screening was also demonstrated by quantitatively recording the positive chronotropic effect of isoproterenol. Thanks to its sensitive signal acquisition ability and unique feature on flexibility, OECTs show a good potential to be a novel in vitro monitoring platform serving both conventional tasks and emerging needs.
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Adv. Healthcare Mater. 2014, DOI: 10.1002/adhm.201400406
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COMMUNICATION Figure 4. a) Photographs of a flexible OECT array device bent to both sides. b) Transfer characteristics of a flexible OECT measured at different bending status in cell culture medium. Negligible change was observed among different bending status. Representative traces of spikes (c) and corresponding zoomed single spikes (d) recorded on the flexible OECT array.
Supporting Information Supporting Information is available from the Wiley Online Library or from the author.
Acknowledgements C.Y. and Q.L. contributed equally to this work. The authors thank financial support from the Theme-based Research Scheme from Research Grant Council (RGC) of the Hong Kong SAR Government, China (Project number: T13–706/11–2). Received: July 14, 2014 Revised: September 25, 2014 Published online:
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