- Record: found
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High-performance transistors for bioelectronics through tuning of channel thickness
research-article
Author(s):
Jonathan Rivnay
1 ,
Pierre Leleux
1
,
2 ,
Marc Ferro
1 ,
Michele Sessolo
1 ,
Adam Williamson
3
,
4 ,
Dimitrios A. Koutsouras
1 ,
Dion Khodagholy
1 ,
Marc Ramuz
1 ,
Xenofon Strakosas
1 ,
Roisin M. Owens
1 ,
Christian Benar
3
,
4 ,
Jean-Michel Badier
3
,
4 ,
Christophe Bernard
3
,
4 ,
George G. Malliaras
1
,
‡
Publication date (Electronic):
22 May 2015
Journal:
Science Advances
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Abstract
Transistors with tunable transconductance allow high-quality recordings of human brain rhythms.
Abstract
Despite recent interest in organic electrochemical transistors (OECTs), sparked by their straightforward fabrication and high performance, the fundamental mechanism behind their operation remains largely unexplored. OECTs use an electrolyte in direct contact with a polymer channel as part of their device structure. Hence, they offer facile integration with biological milieux and are currently used as amplifying transducers for bioelectronics. Ion exchange between electrolyte and channel is believed to take place in OECTs, although the extent of this process and its impact on device characteristics are still unknown. We show that the uptake of ions from an electrolyte into a film of poly(3,4-ethylenedioxythiophene) doped with polystyrene sulfonate (PEDOT:PSS) leads to a purely volumetric capacitance of 39 F/cm 3. This results in a dependence of the transconductance on channel thickness, a new degree of freedom that we exploit to demonstrate high-quality recordings of human brain rhythms. Our results bring to the forefront a transistor class in which performance can be tuned independently of device footprint and provide guidelines for the design of materials that will lead to state-of-the-art transistor performance.
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Most cited references38
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A general relationship between disorder, aggregation and charge transport in conjugated polymers.
Rodrigo Noriega, Jonathan Rivnay, Koen Vandewal … (2013)
Conjugated polymer chains have many degrees of conformational freedom and interact weakly with each other, resulting in complex microstructures in the solid state. Understanding charge transport in such systems, which have amorphous and ordered phases exhibiting varying degrees of order, has proved difficult owing to the contribution of electronic processes at various length scales. The growing technological appeal of these semiconductors makes such fundamental knowledge extremely important for materials and process design. We propose a unified model of how charge carriers travel in conjugated polymer films. We show that in high-molecular-weight semiconducting polymers the limiting charge transport step is trapping caused by lattice disorder, and that short-range intermolecular aggregation is sufficient for efficient long-range charge transport. This generalization explains the seemingly contradicting high performance of recently reported, poorly ordered polymers and suggests molecular design strategies to further improve the performance of future generations of organic electronic materials.
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In vivo recordings of brain activity using organic transistors
Dion Khodagholy, Thomas Doublet, Pascale Quilichini … (2013)
Most breakthroughs in our understanding of the basic mechanisms of information processing in the brain have been obtained by means of recordings from electrodes implanted into, or placed on the surface of the brain. Such recordings allowed the discovery of place cells, grid cells, mirror neurons1 2 3 and more. They also provided an insight into the organization of the brain4, and showed that oscillations constitute the hallmark of brain activity. These oscillations are divided into different frequency bands, from ultraslow to ultrafast, including delta (0.5–3 Hz), theta (4–12 Hz), gamma (40–80 Hz), ripples (100–200 Hz) and sleep spindles (>500 ms long 10–14 Hz oscillations)4 5. Specific oscillations are also recorded in pathological contexts such as spike and wave discharges (SWDs) between 7 and 11 Hz in experimental models of absence epilepsy. For other types of epilepsies, the epileptogenic regions may be determined by assessing the presence of interictal spikes and/or very fast ripples (>200 Hz)6. State-of-the-art recordings are currently performed with microfabricated arrays of metal electrodes (silicon probes, Utah arrays and tetrodes7), which capture the local field potentials (LFPs) generated by the spatio-temporal summation of current sources and sinks (caused by the flux of ions through ion channels localized in the cell membrane) in a given brain volume4. Such probes are also being used in the clinic to improve diagnosis and treatments. For example, stereoelectroencephalography and electrocorticography (ECoG) probes are used to localize epileptogenic zones and to provide functional mapping of the brain before surgery8 9 10 11 12. Although ECoG probes are easier to use than stereoelectroencephalography probes, recordings performed on the brain surface pick up a highly integrated, global signal, which corresponds to the summation of different signals generated at different depths. Hence, ECoG probes are not able to accurately detect activities generated by smaller cell assemblies, except those generated right below an electrode. In addition to clinical applications, microfabricated probes are also likely to have a key role in the design of future brain-machine interfaces13 14. However, major technological advances are still needed; the probes must be fully biocompatible (to enable long-term recordings), small/thin (to decrease invasiveness), highly conformable (to comply with the complex three-dimensional architecture of the brain) and, most importantly, must provide an increased signal-to-noise ratio (SNR) through a built-in pre-amplification/processing system. Neurons and brain networks generate small electric potentials, which are difficult to extract from noise when recorded with classical electrodes made of metals such as Ir, Pt and Au. Advances in microelectronics have given rise to the electrolyte/oxide/silicon field-effect transistor (FET), a more sophisticated device that has been used to measure in vitro signals from cell cultures and tissues slices15 16 17 18. In these devices, the transmembrane current from a neuron in the electrolyte polarizes the gate dielectric and leads to a change in the conductance of the underlying silicon channel. The use of transistors rather than simple electrodes provides the potential of increased SNR due to local amplification endowed by the transistor circuit, and of massive integration, which is possible through the use of matrix-addressing technology developed for flat-panel displays15 16 17 18. These advances, however, have so far been limited to in vitro recordings, mostly because of the poor biocompatibility of the oxide layer of the FETs. Although silicon FETs have recently been integrated into in vivo probes as a means of enabling simultaneous addressing of hundreds of electrodes19, the recordings were carried out by classical electrodes, whereas the transistors themselves were carefully encapsulated to avoid direct contact with the brain. An alternative transistor architecture, termed the organic electrochemical transistor (OECT), was developed in the ‘80s (ref. 20). In contrast to FETs, where an oxide separates the channel from the electrolyte and prohibits any ion transport between these two layers, the channel of OECTs is in direct contact with an electrolyte. As a result, the channel/electrolyte interface constitutes an integral part of the operation mechanism of OECTs. State-of-the-art OECTs are based on the conducting polymer poly(3,4-ethylenedioxythiophene) doped with polystyrene sulfonate (PEDOT:PSS)21. This material is a heavily doped p-type organic semiconductor, in which holes on the PEDOT chains (the semiconductor) are compensated by sulfonate anions on the PSS (the dopant)22. The application of a positive bias on a gate electrode immersed in the electrolyte causes cation injection into the PEDOT:PSS film. These cations compensate the sulfonate anions and dedope the PEDOT, thereby decreasing the drained current23. PEDOT:PSS OECTs work, therefore, in the depletion mode. Accumulation-mode OECTs, based on intrinsic organic semiconductors, have also been reported, but they typically show higher operation voltages24. As OECTs capture ion fluxes25 26, they constitute the optimal solution to measure electrophysiological signals—fluctuations of the electric field (field potentials), generated by the movement of ions27. OECTs offer additional advantages that make them attractive candidates for neural interfacing, including cytocompatibility and straightforward integration with mechanically flexible (hence conformable) substrates17 18. Here we demonstrate the first in vivo use of a transistor to record brain activity. We fabricated highly conformable arrays of OECTs and used them to carry out ECoG on the somatosensory cortex of rats. Simultaneous recordings from penetrating and surface electrodes were used to validate the transistor recordings in two animal models. Compared with surface electrodes, OECTs showed a superior SNR due to local amplification. They also revealed a richer electrophysiological signal, similar to that obtained with penetrating electrodes. Results Structure of the transistor arrays We fabricated ECoG probes that contained OECTs, as well as electrodes made from PEDOT:PSS. Micrographs of an ECoG probe and the layouts of a transistor and an electrode are shown in Fig. 1. A 2-μm thick parylene film was used as the substrate, onto which Au and PEDOT:PSS films were photolithographically patterned. Au served as source and drain electrodes, electrode pads and interconnect lines, whereas PEDOT:PSS was used for the transistor channel and the surface electrodes. A second 2-μm thick parylene film, appropriately patterned to allow access to the channel and to the electrodes, was deposited on top and used as the insulator (Fig. 1c). The total thickness of the arrays was ~4 μm, resulting in probes that were highly conformable yet had enough mechanical strength to be self-supporting and allow manipulation during surgery. Each probe contained 17 transistors with a channel length of 6 μm and a channel width of 15 μm, and 8 electrodes with dimensions of 12 × 12 μm2. The Au structures were completely covered with PEDOT:PSS or parylene, and were not exposed to the electrolyte. Parylene is a Food and Drug Administration-approved polymer used in implantable devices, such as pacemakers, whereas PEDOT-based electrodes have been extensively used as recording electrodes in vivo 28, and have been shown to outperform traditional metal electrodes in chronic experiments29. Implanted in the brain, both parylene and PEDOT:PSS elicit a very small glial response (Supplementary Fig. S1). Therefore, the arrays exposed biocompatible-only materials to the brain. A through hole at the centre of the array allowed the insertion of a silicon probe. At the other end of the probe, pads compatible with a zero insertion force connector allowed easy interfacing to electronics for recording, as shown in the inset of Fig. 1a. In vitro characterization The transistors were characterized in vitro using Ringer’s solution as the electrolyte and a stainless-steel gate electrode. Their output characteristics, shown in Fig. 2a for a drain voltage (V D) between 0 and −0.5 V and a gate voltage (V G) between 0 and 0.5 V, are typical for operation in the depletion regime23. Upon application of a positive gate voltage, cations from the electrolyte enter the polymer film and dedope it, decreasing the drain current (I D). Owing to the absence of a gate oxide and to the high conductivity of PEDOT:PSS, the transistors show a low-operation voltage, which permits operation in aqueous environments. The corresponding transfer curve for V D=−0.4 V, shown in Fig. 2b, exhibits a slope that increases with gate voltage. This is reflected in a transconductance that increases with V G up to 0.42 V, where it reaches a maximum of 900 μS. Normalized for channel width, this value (60 × 103 μS mm−1) is two orders of magnitude larger than that of planar silicon-based FETs used in in vitro neural interfaces16 and three orders of magnitude larger than that of typical organic FETs30, which reflects the efficient gating of the polymer channel due to direct contact with the electrolyte. The transconductance was constant up to 1 kHz (Supplementary Fig. S2), which is above the fastest oscillations recorded in the brain31. It should be noted that the steady-state gate current was less than 10 nA for V D=−0.4 V and V G=0.5 V. Device statistics and stability data are shown in Supplementary Figs S3–S5. In vivo characterization As the first application of OECT-based ECoGs is likely to be for epilepsy diagnosis and cortical mapping, we first characterized the in vivo performance of the transistors in an experimental model of epileptiform activity in rats. Animals were deeply anaesthetized and a craniotomy was performed. The ECoG probe was placed on the somatosensory cortex, and a silicon probe displaying a linear array of Ir electrodes was implanted through the hole in the centre of the ECoG probe (Fig. 3a). The transistor was wired in a common source configuration (Fig. 3b), with the grounded screw used as the gate electrode. We used V D=−0.4 V and V G=0.3 V. Bicuculline, a GABAA receptor antagonist, was perfused on the surface of the brain. Blockade of GABAergic inhibition invariably leads to the genesis of spikes that resemble interictal spikes32. Representative recordings from an OECT, a PEDOT:PSS surface electrode, and the penetrating electrodes of the silicon probe are shown in Fig. 3c. The temporal coincidence of the peaks in the data indicates that the transistor records the same information as the electrodes. The background activity signal is shown at the same scale for the three recording devices, demonstrating the far superior SNR of the transistor. The SNR was calculated by taking the highest peak during a period of epileptiform activity and the standard deviation (STD) of the background during a period of low biological activity. For the OECT recordings, these values were 1.5 μA and 9.5 nA, respectively, yielding an SNR of 44 dB, whereas the PEDOT:PSS surface electrode yielded an SNR of 24.2 dB (4.3 mV peak, 0.26 mV STD background). Although the OECT and the surface electrode were next to one another, and thus picked up the same activity in terms of fluxes of charges on the surface of the brain, the transistor recorded with a much higher SNR. The biological origin of this signal was confirmed by performing current source density (CSD) analysis on the silicon probe recordings (Fig. 3d). The latter reveals the presence of the source and sink (hence, the dipole) generating the epileptiform spike. It should be noted that all recordings were bandpass filtered between 0.1 and 200 Hz (see Methods), to minimize the influence of the acquisition system on the SNR. Although this does not completely eliminate the influence of the acquisition system, it still demonstrates the superior gain provided by the local amplification of the signal by the transistor circuit of Fig. 3b. In the above experiments, the evaluation of the OECT performance was obtained by triggering epileptiform activity with bicuculline. In a second set of experiments, we used a more relevant experimental model, the Genetic Absence Epilepsy Rat from Strasbourg (GAERS)33. This model has been validated in terms of isomorphism, homology and pharmacological predictability to be reminiscent of typical absence epilepsy, a form of generalized non-convulsive epilepsy. GAERS rats show spontaneous large amplitude SWDs at a frequency between 7 and 11 Hz, associated with behavioural arrest and slight perioral automatisms. Despite the fact that deep anaesthesia alters the expression of SWDs in GAERS rats, pathological epileptiform activity could be recorded (Fig. 4a) from the OECT, the PEDOT:PSS surface electrode and from Ir-penetrating electrodes of a silicon probe implanted in the first three superficial layers of the somatosensory cortex. The transistor is shown again to outperform the surface electrode: the SNR for the OECT was 52.7 dB (1.3 μA peak, 3 nA STD background), and 30.2 dB (13 mV peak, 0.4 mV STD background) for the PEDOT:PSS surface electrode. For the sake of completeness, the SNR for the Ir-penetrating electrode was 32.0 dB (10 mV peak, 0.3 mV STD background), although the different location of the probe (depth versus surface recording) does not permit a direct comparison. The time–frequency analysis of epileptiform activity during a short period revealed the presence of typical oscillations around 5Hz (Fig. 4b). Moreover, for better illustration of the SNR improvement the corresponding time traces were normalized according to the peaks of the activity. Electrodes and OECTs alike detected the same signal, in keeping with the fact that SWDs are generalized discharges, which sum up to give rise to a strong signal on the surface. The situation was, however, remarkably different during recordings in between epileptiform activities, where low-amplitude oscillations in the 4–14 Hz frequency range were observed. Time–frequency analysis of recordings showed that the OECT and the Ir-penetrating electrode were able to pick up these low-amplitude signals, whereas the PEDOT:PSS surface electrode showed poor resolution (Fig. 5). The SNR was 22.3 dB for the OECT (300 nA peak, 23 nA STD background), 13.5 dB for the surface electrode (3.5 mV peak, 0.74 mV STD background) and 18.2 dB for the penetrating electrode (3 mV peak, 0.37 mV STD background), with the caveat that the latter recorded in a different location. It should be noted that the depth of anaesthesia used here prevented us from assessing the presence of faster oscillations (such as gamma oscillations) in this case. Discussion The temporal coincidence of the OECT and electrode recordings confirms that both devices record LFPs. The key difference between a transistor and an electrode, however, lies in the fact that the former is an active device, which, when biased as shown in Fig. 3b, forms a circuit that acts as a voltage-controlled current-source amplifier. A voltage-controlled current-source amplifier transforms voltage modulations at the input loop (gate) to current modulations at the output loop (drain), while amplifying the signal power at the same time. For the case of V D=−0.4 V and V G=+0.3 V used in the in vivo measurements, a small oscillation of 50 mV at gate results in an output current oscillation of 37 μA, as defined by the transconductance. At low frequencies, the resistive component dominates the input impedance of the OECT circuit; hence, the input current is determined by the DC bias and is smaller than 10 nA. Therefore, the input signal power is lower than 0.5 nW (=50 mV × 10 nA), whereas the output signal power is 15 μW (=0.4 V × 37 μA). In contrast, conventional electrode recordings are pre-amplified outside the head of the animal. This means that the leads and connections pick up noise that is pre-amplified as well, thereby decreasing the SNR. In general, a higher SNR translates into a shorter overall recording time to obtain the same information. For example, when recording evoked potentials it is necessary to average many individual signals. It also means that new, previously unobserved features can be recorded. Electrodes placed on the surface of the brain record the LFPs associated with the summation of the electrical activity of neural networks, which can be located very far from the recording site. In contrast, penetrating electrodes provide more local, and thus more precise, information on the activity of small populations of neurons. We have previously reported that PEDOT:PSS ECoG electrodes show a better SNR than classical Au electrodes placed on the surface of the brain34. However, PEDOT:PSS electrodes showed less definition as compared with both OECTs and depth electrodes. The use of OECTs for in vivo recordings thus constitutes a major breakthrough, as they are likely to record small and more local activities. This is particularly important in the field of epilepsy, where identifying zones generating high-frequency oscillations or microseizures is the key for diagnosis6 35. OECTs can also provide detailed information on local information processing when functional mapping of brain regions is performed in the operating room before surgery. The biocompatibility of OECTs and their highly conformable nature make them particularly suited for these applications. Small penetrating electrodes are used to capture single-unit recordings, which represent the activity of a single neuron located in the vicinity of the electrode7. This raises the question of whether an OECT placed on a penetrating probe will be able to record the single-unit activity, and what new information, if any, will be revealed by the higher SNR. These experiments are currently ongoing. Finally, OECTs can help answer basic questions in neuroscience about the coupling between electrical activity and metabolism. To function, the brain needs energy in the form of glucose, which is carried in the blood. A dysfunction in this supply system results in pathological activities. Hypometabolism, for example, is one signature of epileptic regions36. The question of how the brain makes use of glucose in different contexts has never been addressed precisely, because it requires the simultaneous recording of neuronal activity and glucose level at the single neuron scale. PEDOT:PSS OECTs coupled with the redox enzyme glucose oxidase have been shown to make simple yet sensitive glucose sensors37. Their integration with electrodes that probe electrophysiology is rather straightforward. Such multi-modal probes would, for the first time, record electrophysiology and metabolism with high spatial resolution. The impact of such probes would be considerable and widespread, in basic physiology, pathology and even in the clinic to interpret metabolic imaging. Methods Transistor fabrication and characterization The fabrication and in vivo validation of PEDOT:PSS electrodes34 and the patterning of PEDOT:PSS OECTs38 were discussed in previous publications. Here we used an adapted fabrication process that involved the deposition and patterning of parylene, Au and PEDOT:PSS films as follows: Parylene C was deposited using an SCS Labcoater 2 to a thickness of 2 μm (at which thickness parylene films are pinhole-free). These films were patterned with the aid of a 4.6-μm thick layer of AZ9260 (MicroChemicals) photoresist and reactive-ion etching by an O2 plasma (160 W, 50 sccm O2, 15 min) using an Oxford 80 plus. Metal pads and interconnects were patterned by a lift-off process. A photoresist, S1813 (Shipley), was spin-coated on the parylene film at 3,500 r.p.m., baked at 110 °C for 60 s, exposed to UV light (150 mJ cm−2) using a SUSS MJB4 contact aligner, and then developed using MF-26 developer. This was followed by the deposition of 5 nm of titanium and 100 nm of gold using a metal evaporator (Alliance Concept EVA450). Lift-off was performed using 1165 stripper. For the preparation of the PEDOT:PSS films, 20 ml of aqueous dispersion (PH-1000 from H.C. Stark) was mixed with 5 ml of ethylene glycol, 50 μl of dodecyl benzene sulfonic acid and 1 wt% of 3-glycidoxypropyltrimethoxysilane (as a cross-linker), and the resulting dispersion was spin-coated at 650 r.p.m. The films were subsequently baked at 140 °C for 1 h and were immersed in deionized water to remove any excess low-molecular weight compounds. The transistors were characterized in vitro using Ringer’s solution (150 mM sodium, 3 mM potassium, 2 mM calcium, 1 mM magnesium and 10 HEPES/NaOH to adjust the pH to 7.2) as the electrolyte. A stainless-steel screw was immersed in the electrolyte and used as the gate electrode. This was the same type of screw that was used as a gate electrode in the in vivo experiments (see below). A Keithley 2612A dual SourceMeter was used to bias the transistor and record the drain and gate currents. The time delay between sourcing voltages and measuring currents was 100 ms, which was found to be long enough for reaching steady-state. In vivo evaluation All protocols have been approved by the Institutional Animal Care and Use Committee of INSERM. Four ECoG probes were tested, two with Wistar rats and two with GAERS rats. The Wistar rats were obtained from Charles River, MA, and the GAERS rats were obtained from Antoine Depaulis (Grenoble-Institut des Neurosciences, Grenoble, France). Upon receipt, they were maintained under controlled environmental conditions (23 °C, 12 h light/dark cycle). The GAERS rats (females, weight of 188 and 196 g, respectively) were initially anaesthetized with 5% isoflurane (Forene, Abbott, France) in 0.5 l min−1 O2 and maintained under anaesthesia with 2% isoflurane. To record faster oscillations, the depth of the anaesthesia was decreased by reducing the amount of isoflurane to 1.5%. The Wistar rats (males, weight of 505 and 551 g, respectively) were anaesthetized with a ketamine/xylazine mixture (35 and 1 mg kg−1, intramuscularly). Additional doses of ketamine/xylazine (7 and 0.3 mg kg−1, intramuscularly) were given as needed. For surgery, the head of the rat was immobilized in a stereotaxic apparatus. The body temperature was monitored and kept constant at 37.5°C with a heating pad. Two miniature stainless-steel screws were driven into the skull above the cerebellum, and served as ground and reference electrodes, respectively. A 5 × 3 mm2 craniotomy was performed in the right hemisphere above the somatosensory cortex (centred at −4 mm in the anterio-posterior axis and −2 mm in the medio-lateral axis, relative to Bregma). The dura mater was removed and the ECoG array was slowly lowered on the surface of the brain. The surface of the cortex was regularly rinsed with a 0.12 M phosphate buffer (33.76 g NaH2PO4–H2O, 7.72 g NaOH in 1 l bidistilled H2O, pH 7.4). The ECoG probe was accessed through a Molex zero insertion force connector with flat flexible cable to flexible printed circuitry configuration. 40 μl of a 100 μM bicuculline solution (Sigma-Aldrich) were deposited with a micropipette onto the brain surface of the Wistar rats after the dura mater was removed. The recordings (Fig. 3c) were taken after sufficient time (20 min) for the effects of bicuculline to diffuse through the cortical layers. In the Wistar rats, an implantable probe (Neuronexus A1 × 16–3 mm 100–177, with a single, 3-mm long shank containing a linear array of 16 electrodes of 177 μm2 area each, spaced at 100 μm from each other) was inserted through the centre of the ECoG in the cortex to reach a final depth of 2 mm. In the GAERS rats, an implantable probe with small pitch between the electrodes (Neuronexus A1 × 8–3 mm 50–177, with a single, 3-mm long shank containing a linear array of eight electrodes of 177 μm2 area each, spaced at 50 μm from each other) was inserted through the centre of the ECoG array in the cortex to reach a final depth of 1 mm. The penetrating probes were connected to a multi-channel Digital Lynx 10S system (Neuralynx) through a X1 HST headstage (Plexon). The signals were amplified (X1,000), bandpass filtered (0.1 Hz–5 kHz) and acquired continuously at 32 kHz on the 64-channel Neuralynx system (16-bit resolution). Signals from the ECoG PEDOT:PSS electrodes were also acquired by the same amplifier and under the same conditions. A Keithley 2612A dual SourceMeter was used to bias the transistor and record the drain and gate currents continuously at 0.3 kHz. Background levels of the recordings The PEDOT:PSS surface electrode recording in Fig. 3c was filtered at 0.1–200 Hz (plus a 50 Hz notch), and we obtain a background level of 0.26 mV. When we digitally filter this recording with a 1–50 Hz bandpass filter, we obtain a background level of 0.2 mV. The electrode recordings in Fig. 4a were filtered at 0.1–200 Hz (plus a 50 Hz notch), and we obtain background levels of 0.4 and 0.3 mV for the surface electrode and for the penetrating electrode, respectively. When we digitally filter the recordings with a 1–50 Hz bandpass filter, we obtain background levels of 0.2 and 0.1 mV for the surface electrode and for the penetrating electrode, respectively. Post-acquisition data treatment All recordings (from electrodes and OECTs alike) were digitally filtered with a 0.1–200 Hz bandpass filter to minimize the contribution of the frequency response of the two different acquisition systems used, and enable a more fair comparison between OECT and electrode recordings. The recordings were also digitally filtered with a 50Hz notch filter to remove line noise. See Supplementary Discussion for the influence of filtering on the background levels of the recordings. The calculation of the SNR was based on recordings filtered in this manner. The data were analysed using MATLAB (MathWorks). A Gabor wavelet time–frequency analysis was used to determine the frequency content of LFPs. CSD analysis of the simultaneously field potentials recorded with the penetrating probe was used to eliminate volume conduction and localize synaptic currents during epileptic discharges. CSD was computed as the second spatial derivative of the recorded raw LFPs (average of 44 events centred on the trough of the epileptic spikes on the deeper electrode of the penetrating probe). The 44 averaged signals superimposed to the CSD were digitally filtered with a 50-Hz notch filter. Power spectra of the recordings are shown in Supplementary Fig. S6. Histology To assess the biocompatibility of parylene and PEDOT:PSS in the brain, we implanted a probe that consisted of an SU‐8 shank in the brain of a 500 g Long‐Evans rat. The shank was 5 mm long, had a width varying from 130 μm near the insertion tip to 400 μm at the end of the shank and a thickness of 50 μm. It was coated with a 2-μm thick conformal parylene film and, on one of its sides, with a PEDOT:PSS film. Parylene and PEDOT:PSS were deposited as described in the methods section. A 2 × 3 mm2 craniotomy was performed in the right hemisphere above the somatosensory cortex, centred at the stereotaxic value of −3 in the antero‐posterior axis and −2 in the medio‐lateral axis, relative to Bregma. A single-shank silicon probe (Neuronexus A1 × 8–5 mm 50–177) was also implanted 1,000 μm apart in the same rat for comparison. Both probes were implanted to a final depth of 2 mm. After 30 days, the animal was killed and the probes removed from the brain. Tissue preparation and histological labelling were performed as described previously39. The rat was deeply anaesthetized with sodium pentobarbital injection (60 mg kg−1, intraperitoneally) and perfused intracardially with a fixative solution containing 4% paraformaldehyde in 0.12 M sodium phosphate buffer (PB), pH 7.4. The rat received 300 ml of this fixative per 100 g of body weight. After perfusion, the brain was removed from the skull, post‐fixed in the same fixative for 1 h at room temperature and rinsed in 0.12 M PB for 1.5 h. It was then cryoprotected in a solution of 20% sucrose in 0.12M PB overnight at 4 °C, quickly frozen on dry ice and sectioned coronally at 40 μm on a cryostat. Sections were rinsed in 0.01 M PBS, pH 7.4, collected sequentially in tubes containing an ethylene glycol‐based cryoprotective solution and stored at −20 °C until histological processing. Every fifth section was stained with cresyl violet to determine the general histological characteristics of the tissue. Author contributions G.G.M. and C.B. conceived the research; D.K. designed, fabricated and characterized the probe; T.D., D.K. and M.G. did the in vivo experiments and data analysis; M.G., P.L. and E.I. helped with the fabrication and computer interfacing of the probes; P.Q. and A.G. helped with the in vivo experiments and data analysis; T.H. and S.S. contributed to the electrode design. Additional information How to cite this article: Khodagholy, D. et al. In vivo recordings of brain activity using organic transistors. Nat. Commun. 4:1575 doi: 10.1038/ncomms2573 (2013). Supplementary Material Supplementary Information Supplementary Figures S1-S6
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High transconductance organic electrochemical transistors
Dion Khodagholy, Jonathan Rivnay, Michele Sessolo … (2013)
High gain is a requirement for transistors in a broad range of applications. In sensors, for example, high gain is associated with high sensitivity, as a small signal in the transistor input yields a large response in the output. There is a tremendous effort to make these devices from soft materials, where low-cost or high-throughput fabrication is desired or where mechanical compatibility with soft tissue is required. The advantageous properties of organic electronic materials, including versatility in processing and synthetic tunability, have allowed for a variety of applications where organic electronic devices are poised to disruptively overtake existing technologies1. A particular organic transistor configuration that is currently attracting a great deal of attention is the organic electrochemical transistor (OECT). Originally developed by White et al.2, OECTs use an electrolyte as an integral part of their device structure. These devices hold considerable promise for technologies such as medical diagnostics and bioelectronic implants due to improved biological and mechanical compatibility with tissue compared with traditional ‘hard’ electronic materials3 4. OECTs have recently been employed in chemical and biological sensing5, and have also been interfaced with cells to control cell adhesion6, and to monitor cell viability7 and barrier tissue integrity8. OECTs have been fabricated using low-cost printing techniques9, and integrated with natural10 and synthetic fibres11 in powerful demonstrations of the unique form factors that can be achieved with these devices. In addition, OECTs utilizing solid or gel electrolytes have widened the scope of application12, showing promise as printable logic circuits13, and drivers for haptic sensors14 and flat panel display pixels15. In most of these applications, transistors convert a modulation in the gate voltage ΔV G to a modulation in the drain current ΔI D. The figure-of-merit that determines this conversion is the transconductance, defined as g m=ΔI D /ΔV G. The transconductance is the main transistor parameter that governs signal amplification: a simple voltage amplifier, for example, can be built by connecting a resistor R in series with the drain. The voltage amplification (voltage modulation across the resistor versus voltage modulation at the gate) is equal to g m R. Despite their many attractive characteristics, transistors based on organic semiconductors are generally not known for their high transconductance, but are instead dismissed in favour of traditional inorganic semiconductors, and, more recently, oxide and graphene-based devices. In this work, we present mechanically flexible OECTs with a transconductance that exceeds that of all other electrolyte-gated transistors and most solid-state devices made of inorganic and low-dimensional nanowire and carbon-based semiconductors. Results Structure of the transistors The device architecture of the OECTs developed for this work is shown in Fig. 1. The channel consists of a 400-nm thick film of the conducting polymer poly(3,4-ethylenedioxythiophene) doped with poly(styrene sulphonate) (PEDOT:PSS), deposited from a commercially available aqueous dispersion and patterned using photolithography. In this material, PEDOT is a semiconductor, which is degenerately doped p-type by PSS to reach conductivities as high as 1,000 S cm−1. Gold source and drain contacts, also patterned photolithographically, define the channel length. The channel length, L, is varied between 5 and 10 μm, and the width, W, is 10 μm. A 2-μm-thick parylene-C film insulates the contacts from the electrolyte solution, and defines an opening where the channel is in direct contact with the electrolyte (a 100-mM NaCl solution). A Ag/AgCl wire immersed in the electrolyte is used as the gate electrode. Steady-state characteristics Figure 2 shows the output characteristics of a typical OECT (L=5 μm, W=10 μm) with negative bias at the drain, V D, and for gate bias, V G, varying from 0 to +0.5 V. These characteristics show the low voltage operation that is the hallmark of electrolyte-gated transistors. The time delay between sourcing V D and V G and measuring the drain current (I D) was 100 ms, which was found to be significantly longer than the time required for the drain current to reach steady-state. The gate current, also measured after the same delay, was <10 nA for V D=−0.6 V and V G=+0.5 V. The corresponding transfer curve for V D=−0.6 V is shown in Fig. 2. The drain current decreases with gate voltage, consistent with operation in the depletion regime. This behaviour is in agreement with our current understanding of the operation mechanism of OECTs16: when a positive bias is applied at the gate, cations from the electrolyte enter the PEDOT:PSS film and compensate the pendant sulphonate anions on the PSS. This leads to a decrease of the hole density in the PEDOT, as holes extracted at the drain are not re-injected at the source. The end result is the decrease of the drain current seen in Fig. 2. This effect is analogous to compensation doping of a p-type semiconductor upon the implantation of donors, but occurs in OECTs at room temperature and at a low applied bias. The transconductance is shown in Fig. 2 to stay above 1 mS over the entire range of applied gate voltage and to reach a peak value of g m=2.7 mS at V G=0.275 V. Devices of similar channel dimensions show a peak g m=2.2±0.8 mS (N=12). The OECT of Fig. 2 is therefore typical and all subsequent data presented here are for this device, unless otherwise mentioned. The best performing device (L=10 μm, W=10 μm) shows a transconductance of 4.0 mS (Supplementary Fig. S1). Table 1 contains a comparison of different transistor technologies15 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 based on their transconductance. The best PEDOT:PSS-based OECT outperforms all electrolyte- and ionic liquid-gated transistors: it outpaces the highest performing electrolyte-gated graphene transistor by one order of magnitude and silicon transistor by two orders of magnitude. It also performs favourably compared with most solid-state technologies, including, amongst others, oxide-gated graphene27 28 and ZnO25. It is only surpassed by III–V semiconductor bulk devices, which, for large channel widths, can yield transconductance values of 30–50 mS33 34. These devices, which require complex multi-layer fabrication, address a different end of the application’s spectrum. Transconductance is often normalized, depending on the device application or materials family. Normalized to the channel width, the OECT reaches 402 S m−1, a value only eclipsed by solid-state III–V semiconductor devices,30 31 32 33 34 and is on the same order of magnitude as oxide-gated graphene transistors27 28. Finally, when the applied voltage is of importance, normalization with respect to V D is used for comparison. The OECT (6,700 μS V−1) again compares favourably with all other devices. Frequency dependence of the transconductance The frequency response of the device shown in Fig. 2 was obtained by measuring the small-signal transconductance. A 100 mV peak-to-peak oscillation was superimposed on the gate bias, and the transconductance was determined by the amplitude ratio between the drain current oscillations and the corresponding input sine wave. Figure 3 shows that the small-signal transconductance stays close to the DC value up to a frequency of ~1 kHz. This is consistent with the temporal response of the drain current upon the application of a voltage pulse at the gate, which is of the order of 100 μs (Supplementary Fig. S2). This behaviour can be understood by considering the fact that OECTs consist of two circuits16: the ionic one, in which ions are transported between the electrolyte and the channel, and the electronic one, in which holes are transported in the PEDOT:PSS channel between the source and the drain. Accordingly, the response time of an OECT can be limited either by ion transport in the ionic circuit, or by the transit time of holes in the PEDOT:PSS channel. The latter was estimated by driving the OECT with constant gate current, I G, and measuring its response time, which, in this case, is solely dependent on the hole transit time, τ e (ref. 16): This relationship was found to hold in the OECTs presented here (Supplementary Fig. S3), yielding τ e =12.6 μs, a value that indicates that hole transport in the channel is not the limiting factor. This is also consistent with the fact that the response time of the OECTs was found to be dependent on the type of ion in the electrolyte (Supplementary Fig. S4). It should be noted that the hole drift mobility that corresponds to this transit time is 0.05 cm2 V−1 s−1. The temporal response of the ionic circuit was evaluated by measuring the impedance, Z, of the channel/electrolyte interface (Supplementary Fig. S5) and calculating the ionic charge, Q, injected in the channel. The latter was derived as: where I G =V G /|Z|, and V G was taken to be the same sinusoidal excitation used to measure the small-signal transconductance (100 mV peak-to-peak). The injected ionic charge is shown in Fig. 3 as a solid line. It shows the same behaviour as the transconductance, signifying that the frequency response of the OECT is dominated by the process of ion transport between the electrolyte and the channel. The agreement between g m and Q is not accidental: Bernards et al.16 showed that the drain current in an OECT is proportional to the ionic charge injected in the channel. A modulation of the gate voltage will lead to a modulation of this charge, and hence, g m will be proportional to Q. Resistance to mechanical deformation In addition to showing a high transconductance, OECTs display a broad range of practically desirable traits that many of the other technologies cannot achieve, such as ease of fabrication, compatibility with mechanically flexible substrates and resistance to aggressive mechanical deformation. Figure 4 shows an array of devices that was peeled-off from its glass sacrificial substrate, crumpled, and then un-crumpled back to a flat sheet. The peeled-off devices had a thickness of ~4 μm, with the PEDOT:PSS channel and the Au interconnects in the neutral mechanical plane. Characterization before and after crumpling showed that the devices could withstand the trial without a significant change in performance. Figure 4d shows the electrical characteristics of one of the devices in this array (L=10 μm, W=10 μm) as-prepared (black), after peel-off from its glass sacrificial substrate (red), and after crumpling (blue), revealing that the transconductance remains practically unchanged. Over 16 devices tested on three different substrates showed little variation in transconductance and time response when exposed to such a harsh handling (Fig. 4f). Discussion The absence of a dielectric between channel and electrolyte in OECTs allows ions to be injected into the former and gives rise to a response that is determined by the volume of the channel. This is in contrast to electrolyte-gated field-effect transistors, where accumulation of charge at the interface between electrolyte and gate insulator (or semiconducting channel) determines the response. This fact is reflected in the magnitude of the charge injected in the channel, which, at low frequencies, is equal to ~2.5 nC (Fig. 3). This corresponds to 7.8 × 1020 ions. cm−3, a value within the range of the doping level of PEDOT:PSS (ref. 41), indicating that the transistor operates close to limit of complete dedoping. As expected, the transistor performance can be tuned by adjusting the channel volume. Making the channel thinner, for example, can lead to faster response, as dedoping it requires a smaller number of ions. At the same time, the drain current (hence its modulation upon gating) decreases, leading to a lower transconductance. Indeed, a device with a PEDOT:PSS channel (L=5 μm, W=10 μm) that was 100 nm thick had a response time of 37 μs and a transconductance of 1.6 mS (Supplementary Fig. S6). In addition to adjusting channel volume, OECT performance can potentially be further improved through the synthesis of new materials. Conducting polymers with a higher capacity to store charge, for example, would lead to devices with a higher transconductance. The design of such polymers is also the focus of contemporary research on batteries42, a synergy that has the chance to lead to the development of new high-performance materials for electrochemical devices. In terms of response time, little is known about ion transport in conducting polymers. Recent measurements by Stavrinidou et al.43 showed that ion mobility in PEDOT:PSS depends on water uptake: pristine PEDOT:PSS films in contact with aqueous electrolytes were found to swell and support transport of small ions with the same mobility as in water. Cross-linking the films diminished their ability to uptake water and decreased ion mobilities. As the relationships between conducting polymer structure and ion transport properties become more established in the future, synthesis of new materials might yield OECTs that operate at higher frequencies. Finally, beyond geometry and materials, the application of a higher drain voltage increases transconductance, but there is a limit of ~1 V to the voltage that can be applied in an aqueous electrolyte before electrolysis takes place. Some ionic liquids and gels enable higher voltage operation, hence higher performance12. The OECT, when biased as shown in Fig. 1, acts as a transconductance amplifier that converts a voltage modulation at the gate to a modulation of the drain current. Considering a biasing point of V D=−0.6 V and V G=0.275 V for the device of Fig. 2, an increase of the input (gate) voltage of 100 mV will result in a change of 270 μA in the output (drain) current. At the same time, the input current remains below 10 nA, increasing by only 2 nA. Therefore, the input signal power is 100 mV × 2 nA=200 pW and the output signal power is 0.6 V × 270 μA=162 μW, yielding a power amplification of 59 dB. Such high power amplification can be harnessed in biosensing applications, which require operation in an aqueous electrolyte and are operated in a DC mode. Interfacing with electrically active cells and tissues is another potential application for OECTs. Extracellular signals, which are of order of 100 μV, can be easily discernible by the OECT, whose transconductance at the frequency of action potentials (1 kHz) is still sufficiently large to ensure significant power amplification. In all these applications, a significant advantage of OECTs as a first stage amplifier is that the input signal is transduced and amplified right at the interface and is hence not corrupted by noise from external wiring and circuitry. In fact, we recently showed that OECTs can record brain activity in vivo with a higher signal-to-noise ratio than electrodes44. Moreover, the fact that OECTs work even after being mechanically deformed opens up a new world of opportunities for devices ranging from smart bandages to intelligent clothing. Methods Device fabrication The fabrication process, similar to that reported previously45, included the deposition and patterning of parylene, metal and PEDOT:PSS. A parylene-C film was deposited on a glass slide using a SCS Labcoater 2 to a thickness of 2 μm (at which the layer is pinhole-free). This layer of parylene was fixed on the glass slide using 3-(trimethoxysilyl)propyl methacrylate as an adhesion promoter. Avoiding the adhesion promoter allowed the devices to be peeled-off after fabrication. An additional 2 μm thick layer of parylene was used to insulate the metal pads from the electrolyte. It was patterned with AZ9260 photoresist and reactive ion etching by an O2 plasma using an Oxford 80 plus. Metal pads and interconnects were patterned by a lift-off process, using S1813 photoresist, exposed to UV light using a SUSS MBJ4 contact aligner, and developed using MF-26 developer. Five nanometre of titanium and 100 nm of gold were then deposited using a metal evaporator, and metal lift-off was performed using acetone. For the preparation of the PEDOT:PSS films, 20 ml of aqueous dispersion (PH-1000 from Heraeus Clevios GmbH), 1 ml of ethylene glycol, 50 μl of dodecyl benzene sulphonic acid were mixed and sonicated before spin-coating. The films were subsequently baked at 140 °C for 1 h and were immersed in deionized water to remove any excess low molecular weight compounds. Electrical characterization All characterization was done using a solution of 100 mM NaCl in DI water as the electrolyte and a Ag/AgCl wire (Warner Instruments) as the gate electrode. The electrical characteristics of the OECTs were measured with two VA10 transimpedance amplifiers (NPI) and customized LabVIEW software. Impedance measurements were carried out using an Autolab PGSTAT 128N. The OECT channel was used as the working electrode (source and drain connected together), while a bridge electrode was used as the reference and a Pt foil was used as the counter electrode. The electrolyte was the same as for the OECT characterization. A modulation with amplitude of 10 mV was applied, while the bias was varied between 0.4 V and −0.4 V, and found to make no difference in the recorded impedance. Author contributions D.K. and G.G.M. conceived the project. D.K., J.R., M.S. fabricated and tested the devices. D.K., J.R., M.G., and P.L. analysed electrical data. D.K., J.R., M.S., M.G., E.S. and L.H.J. developed working mechanism model. T.H., S.S., R.M.O. and G.G.M. supervised the project. D.K., J.R. and G.G.M. wrote the paper. Additional information How to cite this article: Khodagholy, D. et al. High transconductance organic electrochemical transistors. Nat. Commun. 4:2133 doi: 10.1038/ncomms3133 (2013). Supplementary Material Supplementary Information Supplementary Figures S1-S6
Author and article information
Journal
Journal ID (nlm-ta): Sci Adv
Journal ID (iso-abbrev): Sci Adv
Journal ID (publisher-id): SciAdv
Journal ID (hwp): advances
Title:
Science Advances
Publisher:
American Association for the Advancement of Science
ISSN
(Electronic):
2375-2548
Publication date Collection:
May
2015
Publication date
(Electronic):
22
May
2015
Volume: 1
Issue: 4
Electronic Location Identifier: e1400251
Affiliations
[1
]Department of Bioelectronics, École Nationale Supérieure des Mines, CMP-EMSE, MOC,
13541 Gardanne, France.
Author notes
Article
Publisher ID: 1400251DOI: 10.1126/sciadv.1400251
PMC ID: 4640642
PubMed ID: 26601178
SO-VID: cfe890c0-691e-42d2-a6d6-9704e3f325ee
Copyright © Copyright © 2015, The Authors
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This is an open-access article distributed under the terms of the Creative Commons Attribution-NonCommercial license, which permits use, distribution, and reproduction in any medium, so long as the resultant use is not for commercial advantage and provided the original work is properly cited.
History
Date
received
: 22
December
2014
Date
accepted
: 02
April
2015
Funding
Funded by: OLIMPIA;
Award ID: ID0ELGBG1237
Award Recipient
:
Dimitrios Koutsouras
Funded by: ANR, FRM, Region PACA;
Award ID: ID0E4HBG1238
Award Recipient
:
George Malliaras
Funded by: ANR, FRM, Region PACA;
Award ID: ID0ELQBG1239
Award Recipient
:
Christophe Bernard
Funded by: Marie Curie;
Award ID: ID0E6RBG1240
Award Recipient
:
Jonathan Rivnay
Funded by: Marie Curie;
Award ID: ID0ERTBG1241
Award Recipient
:
Michele Sessolo
Funded by: Simons Foundation;
Award ID: ID0EDVBG1242
Award Recipient
:
Dion Khodagholy
Categories
Subject:
Research Article
Subject:
Research Articles
Subject:
SciAdv r-articles
Subject:
Medicine
Subject:
Molecular Electronics
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