Fig. 6.3  Transition of cyclic voltammograms from reversible to irreversible domain. Case 1: reversible reaction with equal cathodic and anodic peak heights. Case 2: transition from reversible to irreversible reaction with increasing separation between cathodic and anodic peaks. Case 3: irreversible reactions indicated by large separation between the cathodic and anodic peaks. Modified from [18], with permission

called the iR drop or the access voltage, which results from the ohmic losses in the system due to resistance of the electrolyte or tissue. These iR losses do not contribute to the potential difference across the interface that drives the charge across the interface. Hence, before performing analysis of the potential transients, it is essential to subtract these losses from the total voltage response. Based on the net potential across the electrode for a given pulse amplitude and duration, estimates of the actual charge injection capacities can be made. Also, by monitoring the voltage drop across the electrode, the safe charge injection limits can be estimated for voltage drops that do not exceed the water window of the electrode.

6.4  Overview of Electrode Materials for Neural Stimulation

An ideal candidate for electrode material for neural stimulation is one which is biocompatible, mechanically stable to surgical implantation, maintains its electrical and mechanical properties for the entire duration of use, is able to support the charge-injection requirements without inducing damage to itself or to the target tissue. Parameters that govern the efficacy and safety of electrode materials have already been described in the preceding section while parameters that govern the safety of biological tissue will be described in Sect. 6.5. In this section, brief overview of materials that are most commonly used as electrodes for neural stimulation will be provided.

Platinum and its alloys with iridium is the most widely used electrode material for neural stimulation. Being a noble metal, it is highly resistant to corrosion and hence suitable for chronic implantations. The electrochemistry of platinum has been well studied along with its charge storage and injection capacities. Along with double layer charging, charge injection can occur through the reversible adsorption of hydrogen onto the platinum surface (H-atom plating) responsible for the pseudocapacity of platinum. Brummer and Turner studied the underlying mechanisms during charge injection through platinum electrodes and its alloys [6–8] and found that these chemically reversible processes can provide charge injection up to 300–350 mC/cm2 in simulated cerebrospinal fluid [8]. In practice, the safe charge injection limit of platinum depends upon a variety of factors such as the pulse duration, current density and geometry of the electrode surface. For square pulses 0.2 ms in duration, the safe charge injection limit with platinum was found to range from 50 to 150 mC/cm2 [48]. Some studies have attempted to increase the electrochemical safe charge injection limit of platinum by increasing the real surface area of the electrode by roughening and have shown varying degrees of success [26, 57].

Iridium oxide belongs to the category of electrodes that are termed as valence change oxides. The oxide layer can be formed in three different ways. Anodic iridium oxide films (AIROF) are produced through repetitive potential cycling of the bulk metal between 0.0 and 1.5 V vs. a reversible hydrogen electrode in an acid or buffered neutral electrolyte [47]. The activated iridium is highly resistant to dissolution and corrosion and exhibit charge storage capacities ranging from 10 to 240 mC/cm2 [52]. This charge storage capacity depends upon the thickness of the film and even moderate activation can lead to high values. However, during neural stimulation, only a fraction of this charge can actually be used. Weiland et al. found the reversible charge injection limits of AIROF to be about 4 mC/cm2, which is greater than platinum and some other metals used for neural stimulation [55]. Beebe et al. showed charge injection limits of about 2 mC/cm2 for biphasic pulses and 3.5 mC/cm2 for monophasic pulses, 0.2 ms in duration with activated iridium wire electrodes [4]. Iridium oxide films can also be formed by thermal decomposition of layers of iridium salts (TIROF) or by reactively sputtering the oxide films onto a substrate from an iridium target (SIROF). Iridium oxide films on the whole have exhibited poor stability during chronic stimulation regimes however, recent work on SIROF shows improvement in in vitro stability during long-term pulsing [9]. In a separate study, Weiland et al. found the metal-tissue interface to be altered after chronic stimulation using thin film iridium oxide electrodes implanted in guinea pig cortex. They observed that current pulsing within safe limits increased the impedance at low frequencies (<100 Hz) after 1 or 2 days of stimulation and found the impedance change to correspond to a reduction in the charge storage capacity [54]. Other studies have also found iridium oxide electrodes to delaminate under high current pulsing with deposits in the surrounding tissue [10].

Capacitive electrodes are ideal for neural stimulation as they do not involve any reactions for charge injection and hence do not have to deal with problem of irreversible reactions. However, these electrodes still have to be operated within the water window in order to avoid hydrolysis. The metal is insulated from the solution

by a thin layer of dielectric material that must be able to withstand the electric fields without any significant dc leakage. Materials in this group that have been found to be safe are anodized tantalum (Ta/Ta2O5), anodized titanium (Ti/TiO2), thin films of barium titanate (BaTiO3) and sputtered deposited titanium nitride (TiN). While anodized tantalum was found to have higher charge storage capacities than anodized titanium and thin films of barium titanate, titanium nitride was found to have charge storage capacities of 23 mC/cm2 when combined with CMOS technology to develop microcolumnar structures [52]. However, the injectable charge limit of titanium nitride was found to be about 0.87 mC/cm2 for microelectrodes while for Ta/Ta2O5 to be around 0.1–0.2 mC/cm2 for large electrodes [47]. Hence, capacitive electrodes though safer than electrodes employing Faradaic reactions, have in general lower charge injection capabilities when operating within the water window.

Carbon nanotubes are also part of the capacitive electrode category exhibiting interesting electrochemical and mechanical properties. They are about five times stronger than steel and yet can be bent and twisted without breaking them. Recent work has shown them as potential electrode material for neural stimulation. Wang et al. developed vertically aligned multiwalled carbon nanotubes (CNTs) using catalytic thermal vapour deposition system [53]. They tested the properties of the CNTs and found that CNTs have a higher charge injection limit of 1–1.6 mC/cm2 after some surface treatment had been performed. Also, continuous pulsing did not degrade the properties of the CNTs. They also found these carbon nanotubes to be capable of causing neuronal excitation in embryonic rat hippocampal neurons. With its precise control of size, geometry and location by lithographic patterning of the catalyst and high charge injection capabilities without any Faradaic reactions, carbon nanotubes may be an answer to the requirements of neuroprostheses employing localized chronic neural stimulation. However, CNTs generally are formed at very high temperatures, making them incompatible with most batch electrode processes.

Conductive polymers are one of the more recent members to the family of electrode materials for neural stimulation applications. Quite a few recent studies illustrate the feasibility of electrochemically polymerizing polypyrrole, polythiophene and their derivatives from aqueous solutions and depositing them on microelectrodes [12–15, 31, 45, 46, 58]. Some of these studies have also shown that these polymers can successfully be incorporated with cell adhesion molecules, growth factors, etc. to further enhance their properties. With its superior electrochemical stability and biocompatibility, poly (3,4-ethylenedioxythiophene) or more commonly known as PEDOT may be well suited for chronic neural interfaces. Recent work suggests that PEDOT coatings can be deposited over platinum electrodes and be used for chronic neural stimulation [16]. The impedance of PEDOT coated electrodes was found to be lower than the bare platinum electrodes with corresponding lower voltage excursion to applied current pulses in PBS. However, the stability of the PEDOT coated electrodes under chronic stimulation regimes was found to depend largely upon the thickness of the coating that can be controlled through deposition time. Physical degradation and changes in microstructure of the film have been suggested as possible­ modes of failure. Hence, more work needs to be done to make these polymers successful electrode materials for chronic neural stimulation.

6.5  Overview of Extracellular Stimulation

The bilipid layer membrane separates the intracellular region of the cell from the extracellular environment and acts as a barrier to the movement of ions between these two regions. It plays a crucial role in determining which ions are allowed to pass through and hence has the important properties of specificity and ­selectivity. The membrane also includes two specialized regions, the afferent region at which the neuron receives the signal and the efferent region at which the neuron sends the signal.

All cells have a resting transmembrane potential (from hereon referred to as membrane potential) with the interior of the cell negative with respect to the exterior of the cell. This membrane potential is dependent on the concentration of the ionic species such that the equilibrium potential of each ion differs from the membrane potential. In general, the ions of interest are K+ (potassium), Na+ (sodium) and Cl− (chloride). At rest, concentration of K+ ions is higher inside giving it a negative equilibrium potential compared to the membrane potential. This gradient tends to move the ions out of the cell. The concentration of Na+ ions on the other hand is higher outside than inside the cell giving it a positive equilibrium potential, which causes them to move into the cell. At rest, the membrane acts as a barrier and is less permeable to Na+ ions compared to K+ ions. The concentration ratios of these ions are maintained by ionic pumps that force the movement of each of the ions in opposite directions thus maintaining a constant charge separation across the membrane and keeping the cell at its resting membrane potential. A typical value of the resting membrane potential is −60 mV measured inside the cell with reference to the outside. Using (6.1), the membrane potential associated with each of the ions is:

where z is the valence of the ion. Although the membrane potential is dependent upon the ionic fluxes, it is not equal to either of their membrane potentials. Instead, the membrane potential of the cell is determined by the concentrations of the ions inside and outside the cell along with the ease with which each of ions can cross the membrane, i.e. on the conductivity and permeability of the membrane to the specific ions. The Goldman equation describes quantitatively the dependence of the membrane potential at steady state on ionic concentration and permeability (P):

As mentioned previously, the cell membrane is selectively permeable to certain ionic species. This is possible due to the presence of ion channels that are pore-like structures spanning across the membrane. As an example, the potassium channels remain open causing a leak of K+ ions out of cell making the inside of the cell more negative. During neuronal signaling, the membrane potential rapidly changes in

response to some stimulus. This is in part achieved by the reduction of membrane potential (depolarization) that leads to the opening of voltage-gated sodium ion channels causing a further reduction of membrane potential. Initially the cell’s response is proportional to the stimulus strength, i.e. the cell responds as a graded potential. Once the membrane potential crosses threshold, the cell responds by generating an action potential that propagates down the cell’s axon all the way to its axon terminals. The axon terminals in turn connect to other cells (through synapses) thereby activating them and thus initiating a signaling cascade. The action potential is described as an all-or-none phenomenon, i.e., once initiated it will actively propagate down the axon irrespective of the presence of the initial stimulus. Typically, action potentials last for about a millisecond after which the cells return to their resting state through the inactivation (closing) of voltage-gated sodium channels and activation of voltage-gated potassium channels. These two mechanisms have longer time constants compared to sodium activation but work together to bring the cell back to its resting membrane potential. This period of inactivation is called the refractory period.

Electrical stimulation of excitable tissue generates action potentials that in turn initiate neuronal signaling and enable partial restoration of lost functionality in sensory or motor systems. This process requires the extracellular region to be driven more negative by applying a rapid negative charge injection via an extracellular stimulating electrode. For the simplest case of stimulation, a single electrode is placed near the excitable tissue and the electrode is driven as a cathode causing the outside of the membrane to become more negative. This causes the membrane potential to become positive thus leading to a net reduction in the membrane potential (depolarizing the membrane). If on the other hand, the stimulating electrode is driven as an anode, then it will cause the outside of the cell to become more positive than the inside thus causing the membrane potential to become more negative. This will lead to a net increase in the membrane potential causing the membrane to hyperpolarize. Since a current generator must have a source and a sink, during extracellular stimulation, a second electrode is required for the current loop to be complete. This second electrode is usually called the return electrode and based upon its size and position can cause a number of different events to occur. If the return electrode is much larger than the stimulating electrode, then the current density is highest at the stimulating electrode causing excitation of neurons near it. However, if the return electrode is similar in size as the stimulating electrode, then the current density at both sites will be the same and hence neuronal excitation can occur at both sites. In this case, during cathodic stimulation, the neurons in close proximity to the stimulating electrode are depolarized while those underneath the return electrode are hyperpolarized. In some case this hyperpolarization may be large enough to suppress an action potential initiated near the electrode (anodic surround block) [41, 43]. On the other hand, if anodic stimulation is employed then the neurons near the stimulating electrode will be hyperpolarized while those near the return electrode will be depolarized. In this case, the action potentials are initiated in regions distant from the electrode known as virtual cathodes. The depolarization that occurs through anodic stimulation is about a seventh to a

third of that accomplished through cathodic stimulation although this depends upon the electrode position [44]. Thus, cathodic stimulation requires less current to cause a cell to cross threshold and initiate action potentials.

The stimulation protocols described above may be effective at selectively activating one population of neurons without activating neighbouring neurons.

Activation thresholds are usually defined in terms of the amount of current needed to cause the excitation along with the duration that the current is applied. Another way to define excitation thresholds is in terms of the applied charge that is simply a multiplication of the amplitude of the applied pulse (current) with the duration of the pulse. Since in neural stimulation, currents applied are in the range of microamps and are applied typically for a few milliseconds, the charge delivered ranges from a few microcoulombs to a few nanocoulombs. By far, the best known law of stimulation is the one by Lapicque that relates the threshold current (I) required for stimulation to the duration (d) of the applied pulse [32]. He introduced the tissue specific excitability parameter called the chronaxie (c) and defined it as the pulse duration that required twice the rheobase current (b). Here, rheobase ­current is defined as the threshold current (I) for very long pulses. Mathematically, b is the limit of I, as pulse duration goes to infinity. The Lapicque law for stimulation is:

Based on the above equation, strength–duration curves can be plotted to graphically illustrate the relationship between the three parameters I, d and c, as shown in Fig. 6.4. The strength–duration curve is an essential tool in all types of studies where electrical stimulation of excitable tissue is employed. Studies have shown how different parameters can be calculated from these curves including charge and energy–duration relationships [24]. Although numerous studies illustrate chronaxie values of different excitable tissues, the accuracy of the measurements can be affected by factors such as the electrode characteristics, tissue inhomogeneity, stimulus waveform, etc. [22, 23]. Studies in motor nerves and different types of muscle have shown the dependence of chronaxie on different parameters such as temperature and location of electrodes [22].

Another way to define the relationship between stimulus strength and excitation is through amplitude-intensity function, as shown in Fig. 6.5. This is typically­

used where the response is an evoked potential and generates a plot of the stimulus strength at fixed pulse duration against the amplitude of the evoked response. It helps in determination of true threshold by simply extrapolating the curve to intersect the x-axis. Amplitude-intensity functions are useful because neural prostheses typically operate above threshold to provide a range of sensation or activation. Finally for the case of single units, analysis methods such as poststimulus time histograms (PSTHs) are employed that sort the individual spikes based on their latencies. More sophisticated analyses of a mixture of action potentials produced by multiple cells involve grouping the individual spikes based on their individual waveform characteristics.