Christophe Bernard
email: christophe.bernard@univ-amu.fr

Daniel T Simon
email: daniel.simon@liu.se

George G Malliaras
email: malliaras@emse.fr

ORGANIC BIOELECTRONICS FOR INTERFACING WITH THE BRAIN

published or not. The documents may come

Introduction

The term bioelectronics refers to the interfacing of electronics with living systems. The earliest known example is the 18 th century experiment of Galvani, who applied a potential across a frog's leg and made it twitch. Two centuries later, the first artificial pacemaker was implanted in a human. Today, about 600,000 pacemakers are being implanted every year. [START_REF] Desilva | Heart disease[END_REF] Cochlear implants, which provide the sense of hearing by electrical stimulation of hair cells in the cochlea, have been implanted in about 300,000 deaf people. Other implantable devices approved by the US Food and Drug Administration (FDA) include spinal cord stimulators for pain relief, deep brain stimulators for Parkinson's disease, nerve stimulators for assisted breathing, bladder control and epilepsy, and retinal implants for vision. These devices are bioelectronic actuators, and trigger the response of a tissue or an organ through electrical stimulation. The reverse is also possible, using a variety of bioelectronic sensors that provide electrical readout of certain biological parameters. These include devices such as the blood glucose monitor, as well as electrodes for recording the electrical activity of the brain, various nerves and muscles. An example of the latter (still at early stages of development) are brain/machine interfaces, which use cortical electrodes for controlling artificial limbs and exoskeletons directly from the brain. [START_REF] Fitzpatrick | Implantable Electronic Medical Devices[END_REF] The devices listed above are clearly making direct and significant improvements in the span and quality of life of a large number of people, and this is the primary factor attracting many researchers to bioelectronics. A large fraction of contemporary research in the field focuses on removing materials bottlenecks at the biotic/abiotic interface. What is needed is materials that will translate the language of electronics to the language of biology and vice versa, and that will do this while satisfying additional constraints, including biocompatibility and suitable mechanical properties. The language of electronics is simple: it consists of flow of electrons in a solid. The language of biology, on the other hand, is considerably more complicated as we will see below. During the past decade, there has been a great deal of interest in the use of organic electronic materials at the biotic/abiotic interface, and the term, "organic bioelectronics" was coined to refer to this effort. [START_REF] Berggren | [END_REF] One of the key reasons for this trend is the ability of organic materials to provide a more intimate interaction with the biological milieu. One way to see this is in terms of the interaction of ionic and electronic carriers at the biotic/abiotic interface. As illustrated by the cartoons of Fig. 1, this interaction can take place along a surface, or take place throughout the volume of the electronic material. In the latter case, the absence of an interfacial barrier and a high mobility of ions in the electronic material allow a 3D interaction between ionic and electronic carriers and pave the way for efficient information transfer between the two worlds. 4 Fig. 1. Schematics of an inorganic semiconductor, silicon, and an organic semiconductor, PEDOT, at the interface with an electrolyte. The hydrated ion is meant to be the same in both schematics, defining the relative scale. The insets show the action of p-type dopants, boron in silicon and PSS in PEDOT, respectively. [Reproduced with permission from Ref. 4].

A model material for organic bioelectronics is poly(3,4ethylenedioxythiophene) doped with poly(styrene sulfonate) (PEDOT:PSS). PEDOT is a conjugated polymer, hence a semiconductor, and is p-type doped by the sulfonate groups of the PSS (Fig. 1). Its conductivity can be very high (of the order of 1,000 S/cm) [START_REF] Elschner | PEDOT, Principles and Applications of an Intrinsically Conductive Polymer[END_REF] and it is therefore referred to as a conducting polymer. It is commercially available in the form of an aqueous dispersion, from which films can be deposited by traditional solution-based techniques such as spin coating. A crosslinker is usually added to render these films stable towards dissolution and delamination in aqueous environments. Films can also be deposited using vapor phase or electrochemical polymerization, with the latter technique requiring a conducting substrate. In the dispersion, PEDOT and PSS form a polyionic complex. PSS is in excess and forms a shell around a PEDOT:PSS-rich core, yielding colloidal gel particles in suspension. After proper shearing and sonication, such colloidal particles are known to be in the range of 5-30 nm. [START_REF] Elschner | PEDOT, Principles and Applications of an Intrinsically Conductive Polymer[END_REF] In the film these particles are thought to flatten, and interconnect. [START_REF] Elschner | PEDOT, Principles and Applications of an Intrinsically Conductive Polymer[END_REF][START_REF] Nardes | [END_REF] Addition of a co-solvent, such as ethylene glycol, aids in the organization and connection of the PEDOT-rich phase and leads to the formation of a more continuous interconnected network that supports efficient hole transport [START_REF] Okuzaki | presented at the Active-Matrix Flatpanel Displays and Devices (AM-FPD), 2012 19th International Workshop on[END_REF][START_REF] Ashizawa | [END_REF] . At the same time, the PSS phase can support ion transport. Indeed, PSS films are known to support efficient ion transport, 9 thought to occur in water pathways created in the film through water uptake and swelling. The same phenomenon is also seen in PEDOT:PSS films, which hydrate extensively upon contact with water and support high ion mobilities. 10 The biological environment, therefore, becomes part of the structure of a PEDOT:PSS film, allowing a 3D between electronic and ionic carriers.

In this chapter we discuss how ion transport in organic electronic materials paves the way for the conception of bioelectronic devices with high performance or novel characteristics. We focus on simple concepts, and stay above materialsspecific details. We start with a brief introduction of how the brain works, then highlight devices for electrical and biochemical interfacing with neurons.

The human brain

The function of the brain is to provide adequate action in response to internal and external stimuli. For example, when spotting the presence of a predator, it is vital to engage evasive actions. But before the motor command is sent to the muscles, a high level of information processing has already occurred as soon as the predator image was projected on the retina photoreceptors. This information processing is the result of complex interactions between three different basic components of the brain: neurons, glial cells and blood vessels (Fig 2). These three components cannot be dissociated from one another; they form a complex interacting system involved in signal processing. 11 The basic unit for information transfer is molecular. Brain cells are secretory cells that can release various molecules, which can influence the activity of neighboring cells. A prototypical example is neurotransmitters, which are released by neurons on their target cells.

Another key molecule is glucose. It is provided by blood vessels, and can be ultimately transformed into adenosine triphosphate (ATP), the basic fuel of cells. If the brain was only relying upon molecular communication, reaction times would be very slow, as information transfer and processing would depend upon diffusion speed. Neurons are not only secretory cells, they are also electrically active. The molecular signals they receive can be transformed in electrical ones resulting in membrane depolarization of hyperpolarization. If the depolarization reaches a key value, the firing threshold, an action potential is generated. The action potential is a fast (ms range) event, which can travel at high speed along the neuron communication line: its axon. At the points of contact between the axon and target cells (the synapse), the action potential triggers the release of neurotransmitters. This mechanism thus allows neurons to transform the result of the integration of the molecular inputs they receive into an electrical signal that will travel fast before it is transformed again into a molecular signal influencing other cells. Historically, the first signals measured in the brain were electrical, in the form of electric fields, known as field potentials. What is the origin of field potentials? The membrane of neurons includes proteins that enable the exchange of ions between the intracellular milieu and the extracellular space. Ion flow is driven by electro-chemical gradients, established by concentration differences and by the polarization of the cell membrane. These transmembrane proteins, when activated, in effect create a channel through which specific ions can flow. We can distinguish two main classes of such transmembrane proteins: voltage-gated and ligand-gated channels. Voltage-gated channels possess a voltage sensor. When the membrane potential reaches a threshold value, the 3D conformation of the protein changes and the channel is created. Thus, the action potential corresponds to the sequential activation of Na + and K + voltage gated channels. When the Na + channels open, Na + ions flow in the cell following their electro-chemical gradient, thus creating a flux of positive charges. As a result, the inside of the cell becomes more positive, which makes the membrane potential more positive, a.k.a. membrane depolarization. A few hundreds of microseconds later, K + channels open and K + ions flow out of the cell, following their electro-chemical gradient. The flux of positive charges leaving the cell makes the inside of the cell less positive, in effect repolarizing the membrane. Ligand-gated channels require the binding of a specific molecule, like a neurotransmitter, on the extracellular side of the protein, leading to the formation of the channel. The molecular signals received by neurons are thus transformed into flows of ions across the cell's membrane. According to the net direction of positive and negative charges, the membrane potential constantly fluctuates. When it reaches the threshold for Na + channel opening, an action potential can be generated.

The constant flow of ions across the membrane through voltage-and ligandgated channels produces a current and thus a field potential that can be captured by recording devices. The first brain recordings were obtained from the scalp with electroencephalography (EEG). It was thus discovered that the brain continuously generates oscillations in different frequency bands (e.g. theta rhythm during REM sleep). 12 The signal recovered from the scalp is highly integrated, coming from millions of brain cells, and highly filtered by the bone and skin. In addition, the spatial origin of the signal (i.e. where in the brain it is generated) is not known. Despite these shortcomings, EEG recordings considerably expanded our knowledge of brain function. They also played a key role in investigating neurological disorders such as epilepsy, improving diagnosis and the choice of treatments. Recording brain electrical activity is still considered as the choice method to understand brain function. However, it must be stressed that this electrical activity is already an integrated signal, which is first a molecular one. One can distinguish three main levels of analysis: the microscopic level, to record the activity of voltage-or ligand-gated channels, a more integrated one where action potentials are captured, and a macroscopic level corresponding to field potentials.

As mentioned above, electrophysiology enabled major discoveries about brain function. The sequence of action potentials generated by neurons is assumed to reflect neuronal computation associated to a given function. For example, the recording of action potentials in freely moving rats revealed the presence of place cells in the hippocampus, which code for a specific location in space. [START_REF] O'keefe | Organic Bioelectronics for Interfacing with the Brain[END_REF] Their firing frequency increases as the animal gets closer to this specific spatial location. Since neurons are organized in networks, decoding information processing and brain function often requires recording a maximum number of neurons. State-of-the-art devices include silicon probes with multiple (e.g. 32) small (177 µm 2 ) recording sites made of metals such as Pt or Ir linearly spread out (e.g. every 50 µm) along a rigid shaft that can penetrate brain tissue. Such probes can pick up the activity of up to 100 neurons, and the activity of each neuron can be extracted and separated after complex signal processing (Fig. 3). Action potentials are not the sole signals that can inform us about brain function or dysfunction. Field potentials, which can be viewed as an end result of the activity of brain cells, are also informative. For example, transient bouts of fast oscillations in the gamma band (40-80 Hz) can bind different brain regions involved in a specific task. Such recordings can be obtained directly from neuronal networks (e.g. with silicon probes), at the surface of the brain or from the scalp. As one gets further away from the source of the signal, the latter gets more integrated and less rich in information. Invasive recording devices are used in humans for diagnosis purposes. For example, neurosurgery can be the last hope for drug-resistant patients with epilepsy. Candidates must first undergo presurgical evaluation to determine which regions are responsible for seizure genesis. This involves implanting them with depth electrodes in the suspected areas and/or electrode arrays placed (grids) on the surface of the brain. Field potentials are then recorded during 2 weeks. Signal analysis can reveal which regions need to be removed to cure the patient. Grids have also been placed above the brain surface to record its activity 24/7 up to 2 years in patients with epilepsy. An expert system was trained to detect in the EEG signal signs of an incoming seizure, giving the patient a warning signal. 15 Such systems would greatly improve the quality of life of drug-resistant patients (which represent 30% of the cases).

If it is important to record brain activity for clinical applications and basic science, it is equally important to interfere with it. A prototypical case is Parkinson's disease. Some patients are equipped with devices delivering high frequency stimulations in deep brain structures, enabling a better control of the symptoms. Deep brain stimulation has been tried for different neurological disorders, including epilepsy with mitigated results. 16 These devices are rather crude and bulky and there is room for improvement. For example, it would be very useful to record electrical activity and stimulate with the same device. One could then stimulate brain regions when it is needed, i.e. before seizure onset.

Electrical stimulation is not the only way to interfere with brain activity to treat neurological disorders. Pharmacology has always been the main way to treat patients. But drugs have also serious drawbacks. First, they need to cross the blood brain barrier to penetrate the brain (few molecules can). Second, they penetrate all brain regions, thus affecting all functions, often resulting in deleterious side effects. Third, they can be very toxic outside the brain (e.g. in the liver). One ideal solution would be to deliver drugs only in the regions that need to be controlled. This requires invasive devices, which share the same limitations as recording devices. An ideal system would deliver a drug on-demand (i.e. when needed) and locally via a biocompatible minimally invasive device.

Interfacing electronics with electrolytes

The communication between an electrode and nearby neurons is mediated by ion flow in the electrolyte and the key issue is to maximize the coupling between an electronic current in the electrode and an ionic current in the electrolyte. Although a rigorous treatment of this problem is beyond the scope of this chapter (this review 14 is a good place to start), here we present a simple model, based on concepts from electrochemistry [START_REF] Bard | Electrochemical Methods: Fundamentals and Applications[END_REF] and electronics [START_REF] Spence | Introductory Circuits[END_REF] , that can provide some valuable insights on the role of capacitance in recording and stimulation.

Let us look first at what happens when an electrode is immersed in an electrolyte. We will limit the discussion to the case of an ideally polarizable electrode (IPE), i.e. an electrode that does not participate in charge transfer with the electrolyte, regardless of the applied voltage. This is an idealized case, as at a sufficiently high applied voltage, any electrode will inject/ extract electrons to/from molecules in the electrolyte. Electrodes made of metals such as Pt, Ir and Au, however, behave as IPEs within a limited range of applied voltage.

Let us imagine two IPEs immersed in an electrolyte (Fig. 4). A small voltage V applied between the two electrodes will cause a redistribution of the ions in the electrolyte, building up negative ionic charge close to the anode and positive ionic charge near the cathode. As this happens, a transient current flows and decays to zero when steady-state is reached. At steady-state, a capacitor has formed at each electrode/electrolyte interface, consisting of a sheet of electronic charge in the metal, compensated by a sheet of ionic charge in the electrolyte (the arrangement of ions near the surface is a bit more complex than a single sheet, and takes place in a compact and a diffuse layer. Readers interested in this are referred to an electrochemistry text [START_REF] Bard | Electrochemical Methods: Fundamentals and Applications[END_REF] ). Such a capacitor is called the double layer capacitor, and has a capacitance per unit area that is ~5 μF/cm 2 . The setup can therefore be modeled by a simple equivalent circuit consisting of two capacitors, representing the two electrode/electrolyte interfaces, connected with each other through a resistor, representing the electrolyte. In a practical configuration, one of the electrodes (recording electrode) is made small in order to gain spatial resolution, while the opposite electrode (reference) is large in comparison. As capacitance is proportional to electrode area, the capacitance of the reference electrode is larger than that of the recording electrode and can be neglected from the equivalent circuit. The latter is thus reduced to a capacitor representing the recording electrode, in series with a resistor representing the electrolyte. This simple circuit can be used to gain some understanding on the influence of the electrode capacitance on electrical recording and stimulation of neural activity. An applied voltage across this circuit will cause voltage drops across the resistor and the capacitor

V= 𝑉 ! + 𝑉 " = 𝐼 • 𝑅 + # " , ( 1
)
where Q is the charge on the capacitor. Using the definition of current

𝐼 = $# $% , (2)
+ + + + - - - - V - + + + + - - - - - - - - - - - - - + + + + + + + + + + Voltage Distance C R C
where t is time. This set of equations can be solved for a particular type of excitation to describes the response of the system. In the case of recording neural activity, let us assume that a neuron produces a time-varying flow of ions in the electrolyte that corresponds to an AC current I. One can show [START_REF] Spence | Introductory Circuits[END_REF] that the voltage induced in the circuit is equal to

𝑉 = 𝐼 • 𝑅 -𝐼 • & '•" , (3)
where 𝑗 ) = -1 and ω is the angular frequency. The first term on the righthand side represents the voltage associated with the neural activity in the electrolyte, while the second term corresponds to an undesired voltage drop across the electrode/electrolyte interface. The corresponding complex impedance is given by

𝑍 = 𝑅 - & '•" , (4)
The voltage drop on the capacitor is therefore determined by an impedance term that is inversely proportional to capacitance, and hence can be minimized by increasing C.

A similar argument can be made for stimulation of neural activity. In this case, a current pulse is usually applied to the circuit and the voltage can be shown [START_REF] Bard | Electrochemical Methods: Fundamentals and Applications[END_REF] to evolve as

𝑉 = 𝐼 • 𝑅 + 𝐼 • % " .
(

) 5
The first term on the right-hand side represents the effect of stimulation in the electrolyte, in terms of a voltage that follows the shape of the applied current pulse. The second term corresponds to an overshoot voltage across the capacitor that increases with time, t. This is an undesired effect, as it can cause electrochemical reactions at the electrode/electrolyte interface, and can be minimized by increasing C.

The analysis presented above, although very simple, highlights the need for maximizing electrode capacitance, for both recording and stimulation of neural activity. At the same time, the need for spatial sensitivity necessitates the use of electrodes with a small geometrical area. The key challenge in electrode design, therefore, is to achieve a high capacitance per unit area.

Organic devices that use ions as part of their operation

Conducting polymer electrodes

Small electrodes are required to obtain high resolution recordings and localized electrical stimulation. Decreasing area, however, decreases capacitance (increases impedance), and makes it difficult to record and stimulate effectively. Conducting polymers have emerged as a solution to this problem: they show a lower impedance and better in vivo recording and stimulation characteristics than traditional electrodes. [START_REF] Venkatraman | [END_REF]20 Here we discuss the origin of the low impedance and show contemporary examples of conducting polymers used as neural electrodes.

Conducting polymers attracted interest as a way to bridge the mechanical property mismatch between metal electrodes and brain tissue. [START_REF] Kim | Indwelling Neural Implants: Strategies for Contending with the In-Vivo Environment[END_REF] The bulk of the early work concentrated on polypyrrole, synthesized electrochemically on metal electrodes, but PEDOT emerged as a more stable material. [START_REF] Asplund | [END_REF] It was shown that the modification of metal electrodes with conducting polymers reduces impedance, a fact attributed partly to a rougher surface (increased surface area), and partly to ion uptake in the polymer film. Fig. 5 shows electrochemical impedance spectra from a Au and a PEDOT:PSS electrode of similar dimensions and roughness (the PEDOT:PSS film was deposited from dispersion and had very low roughness). The line shown through the PEDOT:PSS data is a fit to Eq. ( 4), corresponding to a simple RC circuit, and yielding an effective capacitance per unit area of 2500 μF/cm 2 . This value is 500 times higher than the double layer capacitance of an IPE, implying that ions penetrate into the film. The simultaneous presence of electronic and ionic carriers within the conducting polymer can render the analysis of impedance spectra complicated. 23 In general, equivalent circuits consisting of several elements are used to fit the data, considering the path of an ion moving from the electrolyte to the bulk of the polymer. These elements include a resistor to describe transport in the electrolyte, a capacitor (usually modeled with a constant phase element) in parallel with a charge transfer resistor to model the process of ion injection into the film, and a diffusion element (modeled with a Warburg or a constant phase element) to capture ion diffusion inside the polymer. 24 The complexity of these equivalent circuits makes it difficult to isolate and study the process of ion transport using impedance spectroscopy.

Ion transport, however, can be directly visualized using the "moving front" technique. [25][26][27] Fig. 6 shows a planar junction between an electrolyte and a PEDOT:PSS film. An applied voltage drives cations from the electrolyte into the film, while, at the same time, holes drift away from the junction and are extracted at the other end of the film by a metal electrode. This leads to the formation of a dedoped region in the film, in which injected cations compensate the sulfonate anions. This dedoped region begins at the interface with the electrolyte and, as time goes by, gradually extends towards the opposite end. Its evolution can be easily visualized, as conducting polymers are inherently electrochromic (meaning that their color changes with doping status 28 ). Fig. 6 shows the evolution of the dedoping front, as revealed by light transmission through the film. The drift length ℓ of the cations, which defines the length of the dedoped part of the film, is therefore measured directly. A simple model can be constructed to analyze the data and extract ion mobility in the film. [START_REF] Stavrinidou | [END_REF] Since the dedoped region has a higher resistance than the dedoped one and the electrolyte, it can be assumed that the applied voltage V drops entirely across it. The drift velocity u of the injected cations is given by

𝑢 = 𝜇 • 𝐸 = 𝜇 • * ℓ , ( 6
)
where μ is the ion drift mobility, and E is the electric field across the dedoped part of the film. Substituting

𝑢 = $ℓ $% , (7)
in Eq. ( 6) and integrating over the length of the dedoped region gives

ℓ ) = 2 • 𝜇 • 𝑉 • 𝑡 , ( 8
)
where t is time. The square of the drift length is therefore proportional to time, and the slope gives the ion mobility. Measurements in PEDOT:PSS showed this to be indeed the case for different metal cations, as well as molecular cations such as choline, which were found to be transported with high mobility in the film. 10 It should be stressed, however, that our understanding of ion transport in organic electronic materials is in its infancy, and this topic provides fertile ground for fundamental research. [START_REF] Leger | Iontronics : Ionic carriers in organic electronic materials and devices[END_REF] The low impedance of conducting polymer electrodes, which arises from facile ion transport in the film, allows the fabrication of very small electrodes that can still record local brain activity with a high signal-to-noise ratio (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. A recent example, highlighted in Fig. 7, used arrays of PEDOT:PSS electrodes in an electrocorticography configuration to obtain single neuron activity without penetrating the brain. [START_REF] Khodagholy | [END_REF] The arrays were built on a parylene substrate, onto which Au interconnect lines and PEDOT:PSS electrodes were defined with photolithography, and onto which a second parylene film was deposited to insulate the interconnects. The arrays were only 4 μm thick, providing excellent conformability to the surface of the brain. Fig. 7 shows an array placed on the surface of the brain of an epileptic patient during surgery, and the corresponding recordings of single neuron activity. The low impedance of the electrodes allowed such high quality recordings to be achieved from the surface of the brain, something that could only be accomplished previously by means of penetrating electrodes. These results represent a milestone for the field of organic bioelectronics, and highlight the translational potential of this technology. The low impedance of conducting polymer electrodes also allows efficient stimulation of neural activity. Fig. 8 shows results from pristine and

(a) (b) (c)
PEDOT:PSS-covered PtIr electrodes with a diameter of 25 μm 2 , implanted in rats. [START_REF] Venkatraman | [END_REF] Under biphasic current pulse stimulation, the measured voltage transient corresponding to the PEDOT:PSS is consistent with the one predicted by Eq. ( 5), revealing a lower overshoot than the corresponding traces from the PtIr electrodes. This was correlated with the higher capacitance of the PEDOT:PSS electrodes, measured in vitro.

Organic electrochemical transistors

Signals recorded from the brain are small, and need to be amplified. Noise picked up by the cable connecting the electrode and the amplifier can degrade these signals considerably. Amplifying transducers, which boost up the recordings at the place of acquisition, represent the solution to this problem. Here we show how this can be achieved by leveraging ion transport in conducting polymers, using a device called the organic electrochemical transistor (OECT).

Let us consider the schematic of Fig. 9, which shows a typical setup for recording brain activity with an electrode, and a second one, where recording takes place by an amplifying transducer. The electrode acquires the desired signal S (e.g. a single neuron spike), which is sitting on a background of noise N, caused by the activity of other neurons in the brain, micro-motion of the probe, etc. Additional noise Nline is picked up by the cable leading to the amplifier, owing to electromagnetic noise in the environment. Both signal and noise are then amplified, and recorded. The SNR in this configuration is given by

𝑆𝑁𝑅 = ,•- ,•./,•. !"#$ , ( 9
)
where α is the amplification. In contrast, an amplifying transducer amplifies both signal and noise during the recording, but does not amplify the noise picked up subsequently by cables leading to recording electronics. The SNR in this case is

𝑆𝑁𝑅 = ,•- ,•./. !"#$ , ( 10
)
and, all things being equal, is higher than that of the electrode setup. In practical situations, the noise picked up by cables in a lab can be significant compared to the recorded signals, and a preamplifier is placed as close to the recording site as possible. In in vivo experiments, it is usually mounted at the head of the animal and poses experimental constraints in the design of the probe and its connections. Still, the argument stands, and the ideal solution is to amplify right at the recording site. The latter reflects the conductivity, i.e. the level of doping in the channel. The channel is in contact with an electrolyte, in which a third electrode (gate) is immersed. The gate/electrolyte/channel stack can be modeled as the circuit of Fig. 4. In fact, the capacitance of the gate is usually made large compared to that of the channel so that a voltage is applied at the gate (gate voltage, VG) drops almost entirely across the capacitor defined by the electrolyte/channel interface. The application of VG results in a transient flow of ions towards the channel. What happens next defines the two extreme cases of electrolyte-gated transistors:

In electrolyte-gated field-effect transistors (EGOFETs), the ions accumulate at the surface of the channel and create a polarization, similar to that created at an IPE interface. In OECTs, on the other hand, the ions penetrate into the channel, leading to a large capacitance, as in the case of conducting polymer electrodes. This results in large changes in the doping state of the polymer, meaning large changes in the drain current. OECTs, therefore, are devices that efficiently translate a modulation in gate voltage to a modulation in drain current. It should be noted that it is not presently clear what determines whether ions penetrate an organic semiconductor or accumulate at its surface. Applied voltage, hydrophilicity/hydrophobicity of the material, density, and ion size must play a role in this process, but a more rigorous understanding of how chemical structure affects this process is lacking. PEDOT:PSS, probably due to its hydrophilicity (which translates in the uptake of large amounts of water) is a model system for OECTs. 33

+ Organic semiconductor

Electrolyte

V G + + -- - + + - - --- + + + + + -- + V G + + -- - + + - - - - - + + + + + -- V D V D
Fig. 10. Schematic showing the structure of electrolyte-gated transistors and the difference between an electrolyte-gated field-effect transistor (left) and an organic electrochemical transistor (right).

Fig. 11 shows the transfer curve of a PEDOT:PSS OECT, which indicates how the drain current depends on gate voltage. The data shows that the application of a positive voltage at the gate (which sends cations towards the channel) dedopes the film and decreases the drain current. Similarly, a negative VG will further increase the doping of PEDOT:PSS and increase ID. The slope of the transfer curve is the transconductance, defined as

𝑔 0 = 12 % 1* & . ( 11
)
The transconductance quantifies the "efficiency" with which a transistor transduces a voltage modulation to a current modulation, a transduction that comes with amplification in the power of the signal. Due to the large capacitance associated with ion penetration in their channel, OECTs show a record-high transconductance among all transistors, 33 and as such they make excellent transducers for electrophysiology. It should be noted that the large capacitance associated with the channel of OECTs comes with a penalty to pay in terms of response time. This can be understood by considering a simple circuit consisting of a resistor in series with capacitor, that has a response time of

𝜏 = 𝑅 • 𝐶 . ( 12
)
Signals that are faster than τ will not be able to charge the capacitor, meaning that they will not be able to induce significant changes in the drain current. Brain signals, however, are pretty slow. The fastest activity is a single neuron spike, with a duration of the order of a millisecond. OECTs with suitable response times have been already engineered.

PEDOT:PSS OECTs were used to record brain activity in rats. The OECTs, together with PEDOT:PSS electrodes, were integrated in electrocorticography arrays and placed at the surface of the brain of a GAERS rat, which is a model for absence seizures. Simultaneous recordings from conducting polymer electrodes and OECTs revealed >20 dB increase in the SNR in favor of the transistor during recordings of epileptiform activity (Fig. 12). 34 In a separate experiment, it was shown that the high SNR of OECTs allowed the recording of low-amplitude brain activities, which were poorly resolved with surface electrodes. OECTs therefore are promising amplifying transducers for recording brain activity with a high SNR.

Organic electronic ion pumps

So far we discussed devices that interface with neurons electrically. The ability of organics to transport ions paves the way for devices that do not have an analog in traditional electronics and can enable new modes of communication with cells. We discuss a device called the organic electronic ion pump (OEIP), which utilizes ion transport to deliver biochemical cues to neurons. This functionality enables an entirely different mode of interface with cells and tissue, enabling a level of specificity difficult to achieve with electrical methods alone.

Conducting polymer systems generally require counter ions to balance the charge on the conducting backbone. For example, in the PEDOT:PSS system, the Electrode sulfonate ions on the PSS can balance the positive regions on the PEDOT. In the case that the PEDOT is not charged, other positive ions, e.g. Na + or H + , balance the sulfonate ions on the PSS. This is generally expressed by the simplified chemical equation,

𝑃𝐸𝐷𝑂𝑇 / : 𝑃𝑆𝑆 -+ 𝑁𝑎 / + 𝑒 4 ⇌ 𝑃𝐸𝐷𝑂𝑇 5 + 𝑁𝑎 / : 𝑃𝑆𝑆 4 ( 13
)
where Na + is used as an example of the (mobile) positive ion balancing the sulfonate ions on the PSS. As discussed in the introduction, the PEDOT:PSS film is able to conduct both holes, through the PEDOT-rich phase, and ions, through the PSS phase. This is especially true when the film is soaked in water, forming a hydrogel, and enhancing the ionic conductivity through the PSS phase.

This combined electronic and ionic conductivity can be tailored by patterning regions of a PEDOT:PSS film in a way that degrades or destroys the electronic conductivity while maintaining the ionic conductivity. Specifically, this process, known as overoxidation 35 , degrades the electronically conducting PEDOT backbone by attacking the required C=C double bonds. Indeed, given a thin film of PEDOT:PSS, arbitrary regions of electronically (and ionically) conducting electrodes, and regions of electronically insulating but ionically conducting "ion channels" can be demonstrated. This technique has been used to develop a specialized ionic transport device known as the OEIP 37 (Fig. 13). As shown in the figure, the OEIP consists of a single film of PEDOT:PSS patterned, usually by photolithography and chemical overoxidation, into an anode, an ion transport channel, and a cathode. The ion transport channel (overoxidized PEDOT:PSS) is analogous to the salt bridge of electrochemistry, or the "gel" of conventional electrophoretic tools most commonly associated with genetics analysis, e.g., Western Blot techniques. The OEIP also comprises "source" and "target" electrolyte systems, where the source could be considered a reservoir of the substance to be transported, and the target contains the cells, tissue, etc., which should be effected by the transported substance. Finally, it should be noted that OEIPs are generally operated after being fully soaked in water for several hours, to enhance ionic conductivity.

When a voltage is applied between the anode and cathode, an electric field is established across the ion channel, causing ions from the source electrolyte to be electrophoretically "pumped" toward the target electrolyte. This process is, of course, simple electrophoresis through the ion channel. There are three properties, however, which make the OEIP both unique and useful: (i) the ion channel only allows unidirectional flow of cations (for the case of PSS); (ii) the precise electrochemistry provides extremely high precision "dosage"; and (iii) no liquid is transported, ensuring that the target system is not disrupted.

We can now consider these three properties in more detail. First, in the case of the ion channel being composed of over-oxidized PEDOT:PSS, there is an abundance of fixed negative charge on the PSS. This leads to the situation where cations such as Na + or H + can easily be transported through the PSS. However, anions such as Cl -or OH -are largely excluded from transport since they are electrostatically repelled by all the similar negative charges on the PSS. This effect is known as Donnan exclusion. By this logic, one could exchange the PSS in the channel for a polycation, and thus create an anion transport device which excludes cation transport.

Regarding the second unique property of OEIP substance delivery -precision of the "dosage" -one can consider the electrochemical relationship of Eq. ( 13). Here, you can see that for each electron in involved in the electrical branch of the circuit (e.g., through the power supply and ammeter), there is precisely one ion transported through the channel [Na + in the case of Eq. ( 13)]. In principle, this yields a 1:1 ratio between the electrical current and the flux of transported compounds. This relationship also highlights the temporal response of OEIP delivery: when the electron current stop/starts, so does this ion delivery. Finally, since OEIP delivery and transport is based on electrophoresis through the PSS or some other polyelectrolyte, only ions are transported, not masses of solvent. Furthermore, at the delivery point, this flux of ions builds up a locally high concentration (relative to the target electrolyte), leading to a diffusive gradient with epicenter at the delivery outlet. This results in the third unique property of OEIP delivery: non-flow, diffusive delivery without convective or pressure-based disruption of the (potentially fragile) target system.

OEIPs were initially developed as novel tools for spatio-temporally resolved communication with neural cells on planar surfaces in vitro. Using PSS-based channels, the fundamental signaling ions Na + , K + , and Ca 2+ were delivered and used to trigger Ca 2+ signaling in neuroblastoma cells cultured directly on the OEIP target electrode 37 . Later, an OEIP geometry incorporating a secondary ion channel for pre-filling -and thus faster delivery response -and a miniaturized delivery point down to 10 µm was developed. In these experiments, the neurotransmitter acetylcholine was used to trigger Ca 2+ signals from an outlet on the scale of a single cell and at sub-second temporal resolution. The spatial resolution of OEIP delivery was also demonstrated by the differing responses of cells at 50 and 150 µm from the outlet 38 (Fig. 14). These experiments were the first of several to demonstrate delivery of actual neurotransmitters rather than metal ions. These molecules are small, organic compounds which are vital for neural cell signaling and can cause significant effects in neural tissue even at low concentration. More recently, OEIPs were demonstrated as a novel brain/machine interface in vivo in the modulation of mammalian sensory function. In this case, the device was encapsulated for implantation (Fig. 15), and used to trigger an acute auditory response in the guinea pig cochlea 39 . In the cochlea, there are two populations of sound-transducing hair cells, but only one set of these cells (the inner hair cells) utilized the neurotransmitter glutamate. The OEIP was thus configured to deliver glutamate and thus only effect this population of cells expressing glutamate receptors. Since, indeed, only the inner hair cells responded to the OEIP delivery, these results demonstrated both the biochemical specificity of OEIP technology and its efficacy as an implantable device for influencing neurochemical signaling. OEIP technology can be expanded into a repertoire of "iontronic" circuit components. The OEIPs described above can be considered as an ionic resistor. By incorporating additional layers of cation-or anion-selective materials, and tailoring the electronic/ionic interfaces, recent efforts have succeeded in developing fully chemical, rather than electronic, diodes 40 , transistors 41 , and logic circuits 9,40 . Such active ionic circuitry demonstrates the increasing capabilities and speed of OEIP technology as the device dimensions are minimized. Taken as a whole, OEIP technology thus represents a unique and versatile platform for interfacing the realm of electron-mediated circuits with the realm of chemicalmediated neural circuits.

Outlook

In this chapter we presented the field of organic bioelectronics from the perspective of brain interfacing. We focused on the fact that ion transport in organic semiconductors results in a high capacitance, and showed that this is an advantage for electrodes that record or stimulate brain activity. We also showed that ion transport in organic semiconductors allows the fabrication of devices with a unique mode of operation, and which endow neural interfaces with novel properties.

The process of ion injection and transport from an electrolyte into an organic semiconductor film is not understood at present. This is in part due to the fact that these processes are difficult to study in materials that simultaneously transport electronic charge. To this end, experiments such as the "moving front" technique may provide useful information. The dependence of these processes on chemical structure and film morphology must be understood in order to help guide the synthesis of new, improved materials and the choice of suitable processing techniques.

Our understanding of the physics of devices that utilize mixed (electronic and ionic) conduction, such as the electrochemical transistors and ion pumps discussed here, is at its early stages. While simple models that describe operation in a quantitative manner exist, they need to be rigorously tested and improved to obtain a quantitative understanding. This is important in order to be able to optimize devices and understand their limitations, but also to develop better materials.

In this introductory presentation of the field, we did not discuss issues of biocompatibility and long-term stability, as such issues take a lot longer to elucidate for a new technology and are difficult to relate to fundamental materials properties. The glial scarring caused by probes inserted in the brain arises from a mismatch in mechanical properties of brain and probe, 42 and depends on the size and nature of the entire probe, as well as on the method of insertion and anchoring of the probe to the skull. Soft, flexible probes, with which organic electronic devices are naturally compatible, are known to elicit a smaller foreignbody response. 43 Moreover, compared to metal electrodes on the same probe, conducting polymer electrodes have been shown to be able to record for a longer time. [START_REF] Kim | Indwelling Neural Implants: Strategies for Contending with the In-Vivo Environment[END_REF] By all indications, therefore, organic bioelectronic devices are bound to revolutionize neural interface devices and lead to a better understanding of brain function and dysfunction, and the development of tomorrow's diagnostic and therapeutic tools.

Fig. 2 .

 2 Fig. 2. Scheme of the ménage-à-trois made by glial cells, neurons and blood vessels. Basic energy substrates (glucose and oxygen) are brought by blood vessels. Glucose can be captured by glial cells and neurons, to ultimately produce ATP, the basic fuel of cells. Neurons (and glial cells) can communicate via neurotransmitters, which are released by other neurons and glial cells. Activation of voltage-and ligand-gated channels can lead to ion flux in or out cells. Pumps (such as Na + /K + exchanger) restore homeostasis. [Reproduced with permission from Ref. 11].

14

 14

Fig. 3 .

 3 Fig.3. Model of a local field potential generated by an excitatory synaptic input injected in the distal dendrite of a neutron. The injected current is shown in the inset. Near the local injection site, the field potential is negative (blue), corresponding to a sink of current (blue circle). Near and around the soma, there is a corresponding source of current (orange square). Note that the size,

Fig. 4 .

 4 Fig.4. Schematic showing the steady-state reached when a small voltage is applied between two IPEs immersed in an electrolyte. Double layers formed at each electrode/electrolyte interface define capacitors, which are charged at steady-state. A resistor represents the electrolyte.

Fig. 5 .

 5 Fig. 5. Impedance versus frequency plot for a Au and a PEDOT:PSS electrode. The line is a fit to an RC model.

Fig. 6 .

 6 Fig. 6. The "moving front" technique. (a) Schematic of the planar junction, which allows ion injection (indicated by the arrows) into the polymer only from the side, and detail indicating the charge distribution around the dedoping front (not to scale), (b) evolution of the dedoping front during the injection of potassium cations, where ΔT is the change of transmitted light intensity. The red arrow indicates the drift length of ions at t = 5 s. [Reproduced with permission from Ref. 10].

Fig. 7 .

 7 Fig. 7. Recording of brain activity from the surface of the cortex. (a) Micrograph of PEDOT:PSS electrodes on an ultra-conformable array, (b) surgical placement on the surface of human brain (open circle), with acquisition electronics suspended above the cortical surface, (c) recordings containing spiking activity (scale bars: 20 ms × 40 μV) and sample spike waveforms scale bars (scale bars: 1 ms × 40 μV). [Reproduced with permission from Refs. 20, 31].

Fig. 8 .

 8 Fig. 8. Voltage transients recorded on PEDOT:PSS and PtIr electrodes two weeks after implantation in rats after biphasic current stimulation. The conducting polymer electrodes show a smaller voltage transient and a smaller overshoot voltage after the end of the pulse. [Reproduced with permission from Ref. 19].

Fig. 9 .

 9 Fig. 9. Schematic showing signal acquisition with an electrode (top) and with an amplifying transducer (bottom). S stands for signal, N for noise, α for amplification, and the blue triangle represents an amplifier.

Fig. 11 .

 11 Fig. 11. Transfer curve of a PEDOT:PSS OECT.

Fig. 12 .

 12 Fig. 12. Recordings of epileptiform activity (spikes) with a PEDOT:PSS transistor and a PEDOT:PSS electrode. The SNR is 52.7 dB for the transistor recording and 30.2 dB for the electrode recording. [Reproduced with permission from Ref. 34].

Fig. 13 .

 13 Fig. 13. The organic electronic ion pump (OEIP). (a) Side and (b) top view of the device structure. M + is the species "pumped" through the channel channel (e.g., Na + , Ca 2+ , or some positively charged molecule). [Reproduced with permission from Ref. 36].

Fig. 14 .

 14 Fig. 14. The OEIP in vitro. (a) Microscope image of cells (horizontally elongated dark streaks) cultured on OEIP target side. One cell can be seen near approx. 50 µm from the outlet. Another cell can be seen near approx. 150 µm. (b) Ca 2+ response of cells at 50 and 150 µm during pulsed delivery of the neurotransmitter acetylcholine. The numbers next to the voltage pulses are the pulse duration in seconds. Note the longer pulse required to trigger the cell at 150 µm. [Reproduced with permission from Ref. 38].

Fig. 15 .

 15 Fig. 15. The OEIP in vivo. (a) Side and (b) top schematic of an encapsulated OEIP. The electrodes and reservoirs are housed in sealed silicone tubing. The black arrows indicate the flow of cations through and out of the device. The white arrows indicate inflow of compensating cations from the target solution. (c) Schematic of the guinea pig cochlea experiment, with the OEIP mounted on the round window membrane (RWM), a permeable "input" into the perilymph fluid of the cochlea. The purple units indicate the hair cells effected by glutamate delivery. [Reproduced with permission from Ref. 39].

Acknowledgments

This work was supported by the CG13, the Agence Nationale de la Recherche (through MUSIC and POLYPROBE), the Swedish Governmental Agency for Innovation, the Swedish Foundation for Strategic Research, and the European Commission (through OLIMPIA, OrgBio, and EPITARGET). Thanks are due to Jonathan Rivnay, Michele Sessolo, Klas Tybrandt, Karin Larsson, and Dion Khodagholy who contributed data shown in some of the figures.