Transparent, conformable, active multielectrode array using organic electrochemical transistors Wonryung Leea, Dongmin Kima,b, Naoji Matsuhisaa, Masae Nagasea,b, Masaki Sekinoa,b, George G. Malliarasc, Tomoyuki Yokotaa,b, and Takao Someyaa,b,d,e,1 a Department of Electrical Engineering and Information Systems, The University of Tokyo, Tokyo 113-8656, Japan; bExploratory Research for Advanced Technology, Japan Science and Technology Agency, Tokyo 113-8656, Japan; cDepartment of Bioelectronics, Ecole Nationale Supérieure des Mines, 13541 Gardanne, France; dThin-Film Device Laboratory, RIKEN, Saitama 351-0198, Japan; and eCenter for Emergent Matter Science, RIKEN, Saitama 351-0198, Japan
Mechanically flexible active multielectrode arrays (MEA) have been developed for local signal amplification and high spatial resolution. However, their opaqueness limited optical observation and light stimulation during use. Here, we show a transparent, ultraflexible, and active MEA, which consists of transparent organic electrochemical transistors (OECTs) and transparent Au grid wirings. The transparent OECT is made of Au grid electrodes and has shown comparable performance with OECTs with nontransparent electrodes/wirings. The transparent active MEA realizes the spatial mapping of electrocorticogram electrical signals from an optogenetic rat with 1-mm spacing and shows lower light artifacts than noise level. Our active MEA would open up the possibility of precise investigation of a neural network system with direct light stimulation.
|
organic electrochemical transistors thin films optogenetic transparent electrode
|
A
| multielectrode array |
multielectrode array (MEA), which is a device consisting of a 2D array of microelectrodes, is widely used to detect the action potential of neurons and/or muscle cells (1–3). The characterization of neurons with MEA enables one to accurately measure positions of active/inactive cells, propagation of neural signals, and/or networking among multiple neurons, which cannot be measured by single-point detection (4, 5). The applications of MEA can be classified into two groups—namely, in vitro arrays manufactured on glass and/or petri dishes (6, 7) and in vivo arrays made of microwires (8) and/or silicon-based needles (9). The state of the art microneedle array for detecting the action potential of the brain has realized high spatial resolution of 200 μm (10) and high temporal resolution of 60 μs (11) simultaneously. Recently, a mechanically flexible MEA has been developed to realize conformal contact between curved surfaces of the brain and electrodes with minimal invasiveness (3, 8, 12, 13). The flexible MEA has realized simultaneous recording of local field potentials and action potentials from the human cortical surface without penetrating the brain tissues (14). To advance MEA, an active electrode array, in which each passive electrode is coupled with an amplifier, has been examined to improve signal integrity (15–17). Recently, the organic electrochemical transistor (OECT) has been intensively studied as an active component for the purposes of detection and amplification of biosignals. The mixed electronic and ionic transport in the active conducting polymer layer enables extremely high transconductance of ∼1 mS as well as fast response speed of ∼1 kHz (18, 19). Because of the aforementioned properties, EEG (18), electrocorticography (ECoG) (17), and ECG (20) have all been measured using OECTs, with high signal-to-noise ratio (SNR) of at least 54 dB (17). Furthermore, integration of OECTs acting as sensors and organic field effect transistors (OFETs) as a multiplexer was used to form an active electrode array (21). In addition to a function of signal amplification, an active matrix design can significantly reduce the total number of wirings to access each microelectrode, realizing a scalable MEA www.pnas.org/cgi/doi/10.1073/pnas.1703886114
design. Such an active matrix addresser, realized on a flexible substrate, has been well-researched for various active components, such as OFETs, silicon membrane transistors, and nanowire transistors (2, 22–28). However, optical transparency is needed in MEA for observing biotissues using optical microscopy or modulation of ions imaging (29, 30). For this purpose, thin metal electrodes with typical thickness of ∼3–7 nm are used to achieve transparency with transmittance of ∼30–70% (31). Because thin metal cannot achieve high transmittance and high conductance simultaneously, fully transparent conductors are needed. From this viewpoint, indium-tin-oxide (ITO) (32) and graphene (30, 33) were used to realize a highly transparent MEA comprising transparent electrodes and transparent wires. Such a high transparency is needed particularly for MEA with the large number of multiple electrodes, because wirings occupy most of the area. Furthermore, transparent MEA will create a new tool that can stimulate cells and detect signals via both light and electricity. However, a transparent and active MEA has not yet been realized owing to two technical difficulties. First, many transparent conductors and transistors exhibit large changes of electric performances under the illumination of light. Such photosensitivity, which is usually ascribed to the charge traps or band gaps in transparent materials (34, 35), makes it difficult to detect small changes of active potentials of neurons by photoexcitation. Second, the integration of transparent conductors and transparent transistors requires sophisticated engineering of device designs, because transparency and conductivity of transparent conductors have a tradeoff relation (36, 37). The optimized device structure must be designed such that process compatibility may be maintained. This requirement is much more difficult to meet when MEA is manufactured on flexible substrate owing to its low process temperature. Significance We have developed a transparent, ultraflexible, and active multielectrode array (MEA), which consists of transparent organic electrochemical transistors and transparent Au grid wirings. The micropatterned Au grid showed 60% transparency at 475-nm wavelength. The transparent active MEA showed the spatial mapping of electrocorticogram electrical signals from an optogenetic rat with 1-mm spacing and shows lower light artifacts than noise level. Author contributions: W.L., N.M., G.G.M., T.Y., and T.S. designed research; W.L., D.K., M.N., and M.S. performed research; W.L. contributed new reagents/analytic tools; W.L. analyzed data; and W.L. and T.S. wrote the paper. The authors declare no conflict of interest. This article is a PNAS Direct Submission. 1
To whom correspondence should be addressed. Email:
[email protected].
This article contains supporting information online at www.pnas.org/lookup/suppl/doi:10. 1073/pnas.1703886114/-/DCSupplemental.
PNAS Early Edition | 1 of 6
APPLIED PHYSICAL SCIENCES
Edited by John A. Rogers, Northwestern University, Evanston, IL, and approved August 22, 2017 (received for review March 8, 2017)
In this work, we have successfully manufactured a transparent and active MEA on a 1.2-μm-thick parylene substrate. The device consists of transparent OECTs and transparent metal grid wirings, realizing high transparency on the entire device area. The Au grid wiring with linewidth of 3 μm, which is much smaller than the typical size of neurons, exhibits high transparency of 60% and low sheet resistance of 3 Ω/sq on 1.2-μm-thick parylene substrate. The mechanical durability of the Au grid was performed by applying compression strain on 100% prestretched elastomer. The current fluctuation of OECT is suppressed to less than 0.1% when a strong laser beam (150 mW) is illuminated on the channel of transparent OECTs. Furthermore, a 3 × 5 array of OECT has been manufactured with a total thickness of 3 μm, where each OECT exhibits large transconductance (gm) of 1.1 mS and fast response time of 363 μs. Finally, the feasibility of a transparent and active MEA has been shown by mapping evoked response having an amplitude of ≈800 μV at the exact position where the cortical surface of an optogenetic rat is stimulated by a laser beam. Results The transparent MEA consisting of OECTs and Au metal grid wirings is manufactured on a 1.2-μm-thick parylene substrate (Fig. 1A). The total thickness is as small as 3 μm, including the top encapsulation of a 1.2-μm-thick photoresist epoxy (SU-8) layer, resulting in high flexibility and conformability (8, 38). The fabrication process is shown in Methods and SI Appendix, Fig. S1 in detail. The active layer of OECT is inherently transparent owing to the intrinsic transparency of thin poly(2,3-dihydrothieno-1,4-dioxin)-poly(styrenesulfonate) (PEDOT:PSS) active layers, whereas the transparency of wirings and electrodes is realized by Au metal grids (Fig. 1B). Fig. 1C shows the cross-section of a transparent OECT. Fig. 1D shows the circuit diagram that enables selective access to each OECT by applying voltage to the selected line. A photograph of the 3 × 5 OECT array that is conformal onto a rat brain shows that the cerebrovascular of the brain can be recognized clearly through the devices (Fig. 1E). The Au grid electrodes are carefully characterized to ensure optimal transparency and conductivity. Fig. 2A shows the optical microscopic image of transparent and flexible Au grid electrodes. The transparency and sheet resistance depend on the width (w) of the grid and periodicity (p) of the unit (Fig. 2B). The optical and electrical characteristics are investigated by changing p among 12, 15, and 18 μm, while w is fixed at 3 μm, which is the minimal resolution of our lithography system. Fig. 2C shows the measured transparency (Tm) of the Au grids that are laminated on the supporting glass substrate. The opening aperture ratio (Tc) is defined as (p − w)2/p2. After subtracting the transparency of the glass substrate, Tm is linearly proportional to Tc. Because transparency and conductivity of the metal grid have a tradeoff relationship (SI Appendix, Fig. S2), w and p must be optimally designed. Indeed, low sheet resistance of wires is important to minimize the voltage drop in the wire and subsequently, to apply sufficient voltage from the contact pads to the wired OECT channel. In the following experiment, w and p are set to be 3 and 18 μm, respectively. In this design, Tc is 69%, which is compared with that of the ITO MEA for optogenetic applications (32), whereas the sheet resistance is 3 Ω/sq, which is low enough to operate OECTs. In addition to the high transparency and high conductivity mentioned above, another attribute of Au grids is good mechanical durability. To characterize mechanical durability, Au grids with the thickness of 70 nm are formed on 1.2-μm parylene substrates and covered by a 150-nm-thick PEDOT:PSS layer. For comparison, the reference sample, replacing 70-nm-thick Au grids with 70-nm-thick plane ITO, is also prepared. These devices are laminated onto elastomer that is prestretched by 100%. After the lamination, the compression strain of prestretched 2 of 6 | www.pnas.org/cgi/doi/10.1073/pnas.1703886114
Data lines
A
B
Scan lines Su-8 (passivation)
Au grid
Inter connection
PEDOT:PSS
C 2nd passivation Su-8 (1200 nm)
1st passivation for interconnection Parylene (600 nm) Source/Drain Au grid (70 nm) Substrate Parylene (1200 nm) PEDOT:PSS (140 nm)
D
E device
Scan lines
OECT Electrolyte Data lines
Fig. 1. Transparent electrophysiology OECTs array. (A) Image of the transparent electrophysiology OECT array on a parylene substrate. The passivation layer was patterned by SU-8. (B) Magnified image of a transparent OECT single cell. For transparency, the source and drain of OECT were made by a metal grid of Au. (C) A cross-section of the transparent OECT array. (D) Schematic diagram of the OECT array. A certain voltage was applied to the drain line to be measured. The other lines were all connected to the ground of the circuit for preventing cross-talk. The source lines were connected to each current meter. (E) The 3 × 5 transparent electrophysiology array (black dashed square) on the cortical surface of the optogenetic rat. (Scale bar: 1 mm.)
elastomer is slowly released to form wrinkling structures on the surface of the elastomer. Fig. 2D shows the microscopic images of Au grids on elastomer without compression and with compression strain of 30%. The Au grid and ITO with PEDOT:PSS were patterned with dimensions of 1 × 1 cm. Furthermore, a plain 100-nm-thick Au was deposited at the edge of the sheet of dimensions of 1 × 1 cm, with a contact pad of 1 × 0.5 cm. The sheet resistance of the 1 × 1-cm Au grid and ITO was measured under dry conditions using the two probes method. The sheet resistance is measured with increasing compression strain from 0 to 50%. The sheet resistance of the Au grid sample changes from 3 to 7 Ω/sq at the compression strain of 50%, while that of the ITO sample increases from 79 to 378 Ω/sq (Fig. 2E). The cyclic test is then carried out, applying compression strain of 50% (Fig. 2F). The resistance of the Au grid sample increases by 40 Ω/sq, whereas that of the ITO sample increases by 8 kΩ/sq, showing excellent mechanical durability of the Au grids. The transparent OECTs made with the Au grids are prepared, and their I–V characteristics and time response are examined. For comparison, OECTs with nontransparent Au plane films without mesh structures, which are referred to as nontransparent Lee et al.
w
B
C
100
Transparency (%)
A
p
glass
80 60 40 20
Tc = 0,44,62,69
0 300
140 ITO + PEDOT:PSS Au grid + PEDOT:PSS
120 2
100
10
ITO + PEDOT:PSS Au grid + PEDOT:PSS
APPLIED PHYSICAL SCIENCES
Compression 0%
800
F
3
10
R/R0
E Resistance (Ω/sq)
D
400 500 600 700 Wavelength λ (nm)
80 60
1
10
40 20 0
0
10
0
Compression 30%
10 20 30 40 Compression (%)
50
0
20
40 60 80 100 Cycle number (Compression strain 50%)
Fig. 2. Mechanical stability and transparency of Au grid. (A) Microscope image of metal grid on 1.2-μm parylene substrate. (Scale bar: 100 μm.) (B) Image of a metal grid and square grid unit cell parameters. (C) Transmittance spectrum of the metal grid with various calculated transmittance (Tc). (D) A 3D microscope image of a metal grid under (Upper) 0% compression strain and (Lower) 30% compression strain. (E) Resistance change of ITO (70 nm) and Au grid (70 nm) with PEDOT:PSS (150 nm) when compression strain was changed (the number of samples was three). (F) Resistance ratio after cycling test of ITO and Au grid with PEDOT:PSS when the compression strain applied to the film was 50%.
OECTs, are also prepared. The channel width and length are set to be 90 and 60 μm, respectively (SI Appendix, Fig. S3). The fabrication process is detailed in Methods and SI Appendix, Fig. S1. The measurement setup is schematically shown in Fig. 3A, where a gate bias of OECTs is applied via PBS solution and an Ag/AgCl electrode. Fig. 3B shows the transfer curves. The gm values of transparent and nontransparent OECTs (gm = ΔId/ΔVg) are as high as 2.3 and 2.2 mS, respectively, at gate voltage of 0 V. Note that the difference in gm is only 4%. The response times (τ) of the transparent and nontransparent OECTs are measured by applying a gate voltage pulse with a duration of 1 ms (Fig. 3C). τ is evaluated by fitting experimental data with an exponential function. The τ values of transparent and nontransparent OECTs were as low as 97 and 112 μs, respectively, with a gate voltage pulse of 100 mV (Fig. 3B). The difference in τ is only 13%. The cutoff frequency for this device was 50 Hz, with the transconductance of more than 1 mS at 100 Hz (SI Appendix, Fig. S4). This is sufficient to measure the low-frequency ECoG of an optogenetic animal induced by light (39). Moreover, the leakage current of SU-8 was measured with dc bias of 0.6 V (SI Appendix, Fig. S5). The leakage current of SU-8 was low at 8 nA/cm2 after soaking in PBS of 37 °C with the bias of 0.6 V for 24 h, although it showed poor sealing property when used as the passivation layer for reactive materials, such as Mg (40). Also, the gate leakage current was low at 50 nA after soaking in PBS of 37 °C for 24 h (SI Lee et al.
Appendix, Fig. S6). Therefore, the SU-8 passivation is sufficient for short-term surgical applications. A 3 × 5 array of transparent OECTs is fabricated with 1-mm spacing (Fig. 3D). The channel width and length are 70 and 20 μm, respectively (Fig. 3D, Inset). Fig. 3 E and F shows the statistics of the transconductance and the response time for the array. The average transconductance and response time are 1.1 mS and 340 μs, respectively. The transconductance of the array is slightly lower than that of the discrete devices because of the channel dimension and resistance of the interconnection. With the lengths of the data and scan lines of 3 and 5 mm, respectively, and their width of 0.5 mm, the highest resistance of the grid interconnection can be estimated to be ≈48 Ω. Increasing the thickness of the gold layer beyond the 70 nm would allow for a larger and denser array of transparent OECTs (SI Appendix, Fig. S2). In our array design with the simplest structure in which one cell is formed by one OECT, the cross-talk among sensor cells can be avoided by applying appropriate voltage biases (SI Appendix, Figs. S7 and S8), which is confirmed by the circuit simulator (SI Appendix, Figs. S9 and S10) and an experiment (SI Appendix, Fig. S11) for a 3 × 3 OECT array. The switching of our circuit can be controlled not by the gate bias (Vgs) but by the drain bias (Vds). In addition, the load resistance (RL) of the waveform generator/fast measurement unit (WGFMU), which was used for recording of the current, and the wire resistance PNAS Early Edition | 3 of 6
A
B
D
C
E
F
Fig. 3. Electrical characteristic of transparent OECT and OECT array. (A) OECT during electrical measurement. The electrolyte was placed on the channel of OECTs, and Vg was applied using Ag/AgCl from the top of the electrolyte. (B) Transfer characteristics of the transparent and nontransparent OECTs. The channel width and length were 90 and 20 μm, respectively. The maximum transconductance values were 2.2 and 2.3 mS at a −0.7-V source drain and a 0-V gate voltage, respectively. (C) The response time of transparent and nontransparent OECT. The response times were 110 and 120 μs, respectively. (D) The microscope image of a 3 × 5 transparent OECT array and magnified view of the channel (Inset). The channel width and length were 70 and 20 μm, respectively. (Scale bar: D, 500 μm; D, Inset, 30 μm.) (E) The distribution of transconductance in a 3 × 5 transparent OECT array. (F) The distribution of response times in a 3 × 5 transparent OECT array.
(RW) of the Au grid were low at ≈0 and 10 Ω, respectively. The low RL and RW minimize the cross-talk, while the cross-talk was noticeable when high RL and RW were introduced (SI Appendix, Figs. S12 and S13). The data line in the array is chosen by applying drain voltage (−0.6 V) to the selected line, while the other lines are connected to ground (0 V). The cross-talk evaluation is performed by measuring a cell or a data line while a sine wave of different frequencies is applied to each gate voltage and a drain voltage (−0.6 V) is applied to the measured cell. Both the circuit simulation (SI Appendix, Figs. S8–S10) and the experiment (SI Appendix, Fig. S11) showed clear current fluctuation of the sine wave, which has the same frequency as the measured cell. The feasibility of the transparent OECT array has been shown by ECoG recordings on the cortical surface of optogenetic rats (41, 42). The rats express a light-sensitive gene, channelrodopsin-2, in their nerves and their terminals (43). The ECoG signals are induced by a laser beam at the wavelength of 473 nm. To induce localized brain activity, the rat’s brain surface is surgically exposed and targeted by fiber-guided laser stimulation. The 3 × 5 array of transparent OECTs is attached to the exposed surface of the brain and individually measured signals evoked by light stimulation through the device. The conformal contact between the device film and the surface of the brain is achieved by the ultraflexibility of thin films. The laser beam illuminates the exact spot on the channel region of OECTs. Fig. 4A shows the photograph of the transparent OECT laminated on the cortical surface of the rat with the laser 4 of 6 | www.pnas.org/cgi/doi/10.1073/pnas.1703886114
stimulation (40 mW). Owing to high transparency, the reflection of the laser stimulation on the device surface is minimized. Fig. 4 B and C shows the ΔIds/gm under laser stimulations with different excitation intensities of 17, 40, 60, and 110 mW for transparent and nontransparent OECTs, respectively. The result for the transparent OECT exhibits double the amplitude (700 μV) compared with the nontransparent OECT (350 μV) at the same light intensity of 40 mW. The rms values were 38 and 60 μV for transparent OECT and nontransparent OECT, respectively. With increasing the laser power from 40 to 110 mW, the difference in electric signals between transparent and nontransparent OECTs decreases from 200 to 150% because of the saturation of the photoresponsivity at high intensity. To exclude the possibilities of light artifacts, such as the Becquerel effect or light response of active materials, we measured the evoked potentials, with the rms of 0.03 μA, in response to a pair of light pulses (40 mW) with intervals of 4, 8, 12, and 20 ms (Fig. 4D). Because the repolarization of the first neural excitation does not occur within 4 ms, one signal should be measured by a pair of pulses with an interval of 4 ms or less. One peak of the signal is recorded at intervals of 4 ms. With increasing interval, splitting of signals is clearly observed, because the repolarization of the first neural excitation has been completed. These electrical responses associated with repolarization time, which are known as a characteristic of neurons (39), unambiguously prove that the signals are not light artifacts but actual neural activations. It is important to note that the hindrance Lee et al.
A
C
E
F
APPLIED PHYSICAL SCIENCES
D
B
Fig. 4. In vivo evaluation of transparent OECTs with light stimulation using optogenetic rat. (A) Photograph of the transparent OECT on a neuron-concentrated area of the cortical surface of optogenetic mice with blue laser continuous stimulation through optical fiber (500-μm diameter). (Scale bar: 500 μm.) (B) The recorded signal by transparent OECT with light stimulation of a 475-nm wavelength. The rms was 38 μV. (C) The recorded signal by nontransparent OECT with light stimulation of a 475-nm wavelength. The rms was 60 μV. (D) Evoked ECoG potentials in response to pairs of 2-ms light pulses at 60 mW. The rms was 0.03 μA. (E) Photograph of the transparent electrophysiology array in a neuron-concentrated area on the cortical surface of optogenetic mice with blue laser continuous stimulation through optical fiber (500-μm diameter). (Scale bar: 1 mm.) (F) The spatial distribution of the brain signal intensity measured by a 3 × 5 transparent electrophysiology array. The electric potential was calculated using each OECT’s transconductance, which is measured by a Vg sinusoidal signal and Id.
of the artifact by light stimulation is comparable with the peakto-peak noise level of 0.1 μA (SI Appendix, Fig. S14). Nonetheless, the light is targeted directly onto devices. Because light artifacts are a common issue in optogenetic recordings (30), additional discussion can be found later in the text. Using a 3 × 5 array of transparent OECTs with 1-mm2 spatial resolution, mapping of activation potentials is measured. The device is placed on the cortical surface of an optogenetic rat. The laser beam from the optical fiber stimulates the center spot of the array (Fig. 4E). Fig. 4F indicates a strong ECoG signal with an amplitude of 0.96 μA measured by the OECT at the center of the array. The SNR was 13 dB, whereas the peak-to-peak noise was 0.2 μA, and the rms noise level was recorded as 0.02 μA. Signals are also obtained at the upper and lower cells neighboring the central cell owing to the diffusion of laser fibers. Discussion Here, we discuss the scalability of our circuit design. Previously, a 4 × 4 array of OECT had been fabricated, in which each of the OECTs was individually wired (44). Because an N × N array in the individually wired design requires N2 wires, the area is dominated by wirings for a large N. In contrast, our active matrix design can reduce the number of wirings to 2N, realizing scalability for a large N. Note that our array has a simple design without access transistors. However, it does not suffer from cross-talk issues, because Lee et al.
the current fluctuation of OECTs is measured without load impedance. Another type of active matrix design—each cell consisting of one OECT and one access transistor—can reduce the power consumption and the cross-talk induced by RL and RW as well as simplify the readout circuit (21, 25, 26, 45); thus, it is expected that integration of transparent OECT and access transistor will further improve the scalability from the viewpoint of power consumption and improved resolution of measurement. Finally, we would like to emphasize again that the key to realize transparent active MEA is the elimination of light artifacts, which are frequently major problems in optogenetic neural signal recordings. Note that photosensitivity is suppressed in OECT for two reasons. First, the absorption of PEDOT:PSS in OECTs is significantly reduced in the blue-wavelength regime (46), because highly doped holes in PEDOT:PSS compensate for the site of the charge trap. Second, the Becquerel effect, which is inversely proportional to the site impedance, is suppressed by the low site impedance of PEDOT:PSS (47). Indeed, it is reported that a Pt electrode with site impedance of 50 kΩ at 1 kHz could reduce the light artifact (39). Our PEDOT:PSS layer exhibits a lower site impedance of 10 kΩ at 1 kHz and 1,400 μm2 (SI Appendix, Fig. S15) owing to high conductivity and the porous structure (48). In this way, our device showed no detectable light artifacts, even under high-power laser stimulation (110 mW) (SI Appendix, Fig. S14). PNAS Early Edition | 5 of 6
Methods Device Fabrications. Novec 1700 (3M) dissolved in Novec 7100 (3M) was spincoated on a glass wafer as a sacrificial layer. Parylene dix-SR was deposited as an ultrathin substrate of the device to fabricate 1.2-μm-thick films; 70-nmthick Au formed the pattern via the liftoff process. The device was encapsulated by 600 nm of parylene to separate the layer for interconnection; 70-nm-thick Au for interconnection was deposited by the liftoff process. For passivation, the 1.2-μm-thick SU-8(3005) was patterned by photolithography. The solution of PEDOT:PSS with additives (17) was spin-coated on the film at 2,000 rpm (Spincoater 1H-DX2, MIKASA). Finally, the PEDOT:PSS was patterned by dry-etching (49). Characterization. All characterization of the OECTs was performed using a solution of PBS as the electrolyte and an Ag/AgCl wire (Warner Instruments) as the gate of the OECT electrode. The I–V characteristics of the OECTs and the OFETs were measured with a 4155C (Agilent). The WGFMU module of the B1500 (Keysight) measured the drain currents; the other WGFMU module measured the gate voltages of the OECT to determine the time response. A multifunction generator was used to generate the sinusoidal waves.
isoflurane mixed with air, and the brain skull was incised to expose a cortical surface. Stimulation light with a wavelength of 473 nm was delivered from a laser source (COME2-LB473/586/200S; LUCIR). The neural signal was measured by connecting a semiconducting parameter analyzer (the WGFMU module of the B1500; Keysight) to the attached OECT device while optically stimulating the nerve terminal at repetitive rates of 2 Hz. The neural recording from matrix OECT was measured individually with the sampling rate of 1 kHz. The measured data were obtained from a single trace without averaging. The filtration of the noise above 1 kHz was obtained, implemented from the following formula: yn =
1 yn−1 1 ðxn − xn−1 Þ + ts + + ts , f−3dB f−3dB f−3dB f−3dB
[1]
where x, ts, and f−3dB denote the raw data, the interval time, and cutoff frequency, respectively.
In Vivo Evaluation. Animal experiments were conducted in accordance with the Guidelines of the Animal Experiment Committee at the University of Tokyo. The rat [W-Tg(Thy1-COP4-YFP*)4Jfhy; National BioResource Project of the Rat in Japan; male, 12 wk old, 268 g] was anesthetized using 2–2.5%
ACKNOWLEDGMENTS. We thank Marce Ferro, Pierre Leleux, Jonathan Rivnay, Junhyung Kim, Sunghoon Lee, Robert Nawrocki, Sungjoon Park, and Prof. Makoto Takamiya for the discussions and Prof. Hiromu Yawo for supplying the optogenetic rat. W.L. was supported by the Japan Society for the Promotion of Science (JSPS) through the Program for Leading Graduate Schools. N.M. was supported by the Advanced Leading Graduate Course for Photon Science and a JSPS research fellowship for young scientists. This work was supported by Japan Science and Technology Agency Exploratory Research for Advanced Technology Grant JPMJER1105.
1. Rubehn B, Bosman C, Oostenveld R, Fries P, Stieglitz T (2009) A MEMS-based flexible multichannel ECoG-electrode array. J Neural Eng 6:036003. 2. Viventi J, et al. (2011) Flexible, foldable, actively multiplexed, high-density electrode array for mapping brain activity in vivo. Nat Neurosci 14:1599–1605. 3. Minev IR, et al. (2015) Electronic dura mater for long-term multimodal neural interfaces. Science 347:159–163. 4. Sawada M, et al. (2015) Function of the nucleus accumbens in motor control during recovery after spinal cord injury. Science 350:98–101. 5. Ishida A, et al. (2016) Causal link between the cortico-rubral pathway and functional recovery through forced impaired limb use in rats with stroke. J Neurosci 36:455–467. 6. Oka H, Shimono K, Ogawa R, Sugihara H, Taketani M (1999) A new planar multielectrode array for extracellular recording: Application to hippocampal acute slice. J Neurosci Methods 93:61–67. 7. Heuschkel MO, Fejtl M, Raggenbass M, Bertrand D, Renaud P (2002) A threedimensional multi-electrode array for multi-site stimulation and recording in acute brain slices. J Neurosci Methods 114:135–148. 8. Kim D-H, et al. (2010) Dissolvable films of silk fibroin for ultrathin conformal biointegrated electronics. Nat Mater 9:511–517. 9. Buzsáki G (2004) Large-scale recording of neuronal ensembles. Nat Neurosci 7:446–451. 10. Wark HAC, et al. (2013) A new high-density (25 electrodes/mm2) penetrating microelectrode array for recording and stimulating sub-millimeter neuroanatomical structures. J Neural Eng 10:45003. 11. Harrison RR, et al. (2007) A low-power integrated circuit for a wireless 100-electrode neural recording system. IEEE J Solid-State Circuits 42:123–133. 12. Kim D-H, et al. (2011) Epidermal electronics. Science 333:838–843. 13. Liu Y, et al. (2016) Epidermal mechano-acoustic sensing electronics for cardiovascular diagnostics and human-machine interfaces. Sci Adv 2:e1601185. 14. Khodagholy D, et al. (2015) NeuroGrid: Recording action potentials from the surface of the brain. Nat Neurosci 18:310–315. 15. Fuketa H, et al. (2014) 1 μm-thickness ultra-flexible and high electrode-density surface electromyogram measurement sheet with 2 V organic transistors for prosthetic hand control. IEEE Trans Biomed Circuits Syst 8:824–833. 16. Lacour SP, Jones J, Wagner S, Li T, Suo Z (2005) Stretchable interconnects for elastic electronic surfaces. Proc IEEE 93:1459–1467. 17. Khodagholy D, et al. (2013) In vivo recordings of brain activity using organic transistors. Nat Commun 4:1575. 18. Rivnay J, et al. (2015) High-performance transistors for bioelectronics through tuning of channel thickness. Sci Adv 1:e1400251. 19. Khodagholy D, et al. (2013) High transconductance organic electrochemical transistors. Nat Commun 4:2133. 20. Campana A, Cramer T, Simon DT, Berggren M, Biscarini F (2014) Electrocardiographic recording with conformable organic electrochemical transistor fabricated on resorbable bioscaffold. Adv Mater 26:3874–3878. 21. Lee W, et al. (2016) Integration of organic electrochemical and field-effect transistors for ultraflexible, high temporal resolution electrophysiology arrays. Adv Mater 28:9722–9728. 22. Kim J, et al. (2016) A wearable multiplexed silicon nonvolatile memory array using nanocrystal charge confinement. Sci Adv 2:e1501101. 23. Fang H, et al. (2017) Capacitively coupled arrays of multiplexed flexible silicon transistors for long-term cardiac electrophysiology. Nat Biomed Eng 1:38. 24. Yu KJ, et al. (2016) Bioresorbable silicon electronics for transient spatiotemporal mapping of electrical activity from the cerebral cortex. Nat Mater 15:782–791. 25. Someya T, et al. (2004) A large-area, flexible pressure sensor matrix with organic fieldeffect transistors for artificial skin applications. Proc Natl Acad Sci USA 101:9966–9970.
26. Kawahara J, et al. (2013) Flexible active matrix addressed displays manufactured by printing and coating techniques. J Polym Sci B Polym Phys 51:265–271. 27. Takei K, et al. (2010) Nanowire active-matrix circuitry for low-voltage macroscale artificial skin. Nat Mater 9:821–826. 28. Schwartz G, et al. (2013) Flexible polymer transistors with high pressure sensitivity for application in electronic skin and health monitoring. Nat Commun 4:1859. 29. Williamson A, et al. (2015) Localized neuron stimulation with organic electrochemical transistors on delaminating depth probes. Adv Mater 27:4405–4410. 30. Park DW, et al. (2014) Graphene-based carbon-layered electrode array technology for neural imaging and optogenetic applications. Nat Commun 5:5258. 31. Kossoy A, et al. (2015) Optical and structural properties of ultra-thin gold films. Adv Opt Mater 3:71–77. 32. Kwon KY, Sirowatka B, Weber A, Li W (2013) Opto- μECoG array: A hybrid neural interface with transparent μECoG electrode array and integrated LEDs for optogenetics. IEEE Trans Biomed Circuits Syst 7:593–600. 33. Kuzum D, et al. (2014) Transparent and flexible low noise graphene electrodes for simultaneous electrophysiology and neuroimaging. Nat Commun 5:5259. 34. Yao J, et al. (2011) Electrical and photosensitive characteristics of a-IGZO TFTs related to oxygen vacancy. IEEE Trans Electron Dev 58:1121–1126. 35. Milvich J, et al. (2015) Flexible low-voltage organic phototransistors based on air-stable dinaphtho [ 2, 3-b : 2 0, 3 0 -f ] thieno [ 3, 2-b ] thiophene (DNTT). Org Electron 20:63–68. 36. Vosgueritchian M, Lipomi DJ, Bao Z (2012) Highly conductive and transparent PEDOT: PSS films with a fluorosurfactant for stretchable and flexible transparent electrodes. Adv Funct Mater 22:421–428. 37. Wang DH, Kyaw AKK, Gupta V, Bazan GC, Heeger AJ (2013) Enhanced efficiency parameters of solution-processable small-molecule solar cells depending on ito sheet resistance. Adv Energy Mater 3:1161–1165. 38. Kaltenbrunner M, et al. (2013) An ultra-lightweight design for imperceptible plastic electronics. Nature 499:458–463. 39. Richner TJ, et al. (2014) Optogenetic micro-electrocorticography for modulating and localizing cerebral cortex activity. J Neural Eng 11:016010. 40. Fang H, et al. (2016) Ultrathin, transferred layers of thermally grown silicon dioxide as biofluid barriers for biointegrated flexible electronic systems. Proc Natl Acad Sci USA 113:11682–11687. 41. Ji ZG, et al. (2012) Light-evoked somatosensory perception of transgenic rats that express channelrhodopsin-2 in dorsal root ganglion cells. PLoS One 7:e32699. 42. Ishizuka T, Kakuda M, Araki R, Yawo H (2006) Kinetic evaluation of photosensitivity in genetically engineered neurons expressing green algae light-gated channels. Neurosci Res 54:85–94. 43. Boyden ES, Zhang F, Bamberg E, Nagel G, Deisseroth K (2005) Millisecond-timescale, genetically targeted optical control of neural activity. Nat Neurosci 8:1263–1268. 44. Gu X, Yao C, Liu Y, Hsing IM (2016) 16-Channel organic electrochemical transistor array for in vitro conduction mapping of cardiac action potential. Adv Healthc Mater 5:2345–2351. 45. Someya T, et al. (2005) Conformable, flexible, large-area networks of pressure and thermal sensors with organic transistor active matrixes. Proc Natl Acad Sci USA 102:12321–12325. 46. Jeuris K, Groenendaal L, Verheyen H, Louwet F, De Schryver FC (2003) Light stability of 3,4-ethylenedioxythiophene-based derivatives. Synth Met 132:289–295. 47. Khurram A, Seymour JP (2013) Investigation of the photoelectrochemical effect in optoelectrodes and potential uses for implantable electrode characterization. Conf Proc IEEE Eng Med Biol Soc 2013:3032–3035. 48. Ludwig KA, Uram JD, Yang J, Martin DC, Kipke DR (2006) Chronic neural recordings using silicon microelectrode arrays electrochemically deposited with a poly(3,4-ethylenedioxythiophene) (PEDOT) film. J Neural Eng 3:59–70. 49. Braendlein M, et al. (2017) Lactate detection in tumor cell cultures using organic transistor circuits. Adv Mater 29:1605744.
6 of 6 | www.pnas.org/cgi/doi/10.1073/pnas.1703886114
Lee et al.
