There is a growing interest in alternative methods to achieve focal brain stimulation. Invasive techniques, such as deep brain stimulation (DBS) and vagus nerve stimulation (VNS), involve the delivery of a unidirectional current through implanted electrodes. With these techniques, focality is thought to be enhanced by the use of unidirectional, as opposed to bidirectional, current flow and by positioning the cathode proximal to the site of intended maximal stimulation. With noninvasive, transcranial magnetic techniques, imposing a time-varying magnetic field on the scalp has been the principal method for producing focal stimulation in underlying neural tissue.1, 2
Transcranial magnetic stimulation (TMS) relies on the principle of magnetic induction, whereby a time-varying magnetic field induces current flow when in contact with a conductive medium, such as neural tissue. Since the scalp and skull are transparent to the magnetic field, the focality of TMS is principally determined by the geometry and orientation of the magnetic coil. 3–5
Despite its non-invasiveness, TMS has important limitations as a focal stimulation method. The intensity of the magnetic field falls considerably with distance from the coil.6
With standard methods, depolarizing stimulation is limited to a depth from the cortical surface of approximately 2 cm, corresponding roughly to the gray-white matter junction. 7
Second, the energy transfer is extremely inefficient when converting from the time-varying current in the magnetic coil to the orthogonally-oriented magnetic field, and then to the induced electrical field in brain.1
The amperage of the current in the coil (e.g., 10,000 A) is profoundly greater than the peak amperage in brain. Third, the physical limitations of most magnetic coils restrict their capacity to withstand repetitive, high intensity electrical stimuli before heating excessively or otherwise failing. This limitation in coil design effectively caps the maximal intensity of the induced electrical field.8
Fourth, the shape of the electrical waveform induced in brain is a function of the rate of change (first derivative) of the imposed magnetic field. The induced electrical waveform produced by standard magnetic stimulators is complex with multiple positive- and negative-going peaks. Unlike the rectangular pulse traditionally used with electrical stimulation, the peak-induced current with TMS is expressed only instantaneously, and there is uncertainty about the components of the complex waveform that are responsible for its neurobiological effects.
The development of magnetic seizure therapy (MST) illustrates these limitations. In electroconvulsive therapy (ECT), seizures are readily produced with a single train of transcranial electrical pulses delivered by an electrical stimulator. The evidence that electrical dosage and electrode placement strongly impact on the efficacy and cognitive effects of this intervention justified the development of forms of stimulation that offered greater control over intracerebral current paths and current density. Sackeim (1994) originally proposed that this might be accomplished with high intensity magnetic stimulation.9, 10
After 15 years of intensive engineering efforts, it has been shown that MST is capable of reliably eliciting generalized seizures from motor cortex. 11–13
As yet this technique has not demonstrated the capacity to deliver both focal and substantially suprathreshold stimuli to any brain region.
A number of the limitations of TMS can be overcome through use of repetitive transcranial electrical stimulation (rTES). The shape or waveform of electrical stimuli can be easily manipulated and optimized for efficiency in producing specific neurobiological effects. The intensity or amplitude of the electrical stimulus is easily controlled and, in principle, virtually unlimited in magnitude. The capacity to deliver high intensity stimuli means that rTES is capable of producing depolarization in deep brain structures. Furthermore, electrical stimulators have been designed to deliver pulse trains in which the voltage or current of each pulse or the energy delivered over a pulse train is held constant (called constant current stimulation). TES is a form of constant voltage stimulation and current flow in brain is a function of the voltage of the electrical field and local tissue impedance (called constant voltage stimulation). Charge density and charge density per phase are the key determinants of the neurobiological effects of electrical stimulation 14
. Thus it is advantageous that stimulation techniques follow constant current, rather than constant voltage, principles.
Despite these advantages, the use of rTES as either a research tool or treatment intervention has largely been limited to ECT. One reason for this is that the scalp, CSF, and brain have low impedance, while the impedance of the skull is both high and heterogeneous. Under ordinary circumstances, the skull acts as an insulator and, for example, the bulk of externally applied ECT current is shunted through the scalp and does not enter brain 15–19
. To achieve adequate current density in brain many potential applications of rTES would require sufficiently high charge density in peripheral scalp tissue that the stimulation would be painful. However, this limitation of rTES may be overcome by use of general anesthesia for applications involving seizure induction and local anesthesia for non-convulsive rTES.
Of greater concern is the notion that the smearing of the electrical stimulus is unavoidable given the shunting resulting from the high skull impedance, and this undercut any possibility of achieving focality. However, Amassian and colleagues (1987) demonstrated that rTES could achieve a focality comparable in spatial resolution to rTMS, depending on the directionality of current flow, electrode geometry and electrode positioning 20–22
Specifically, one could achieve focality with rTES by using unidirectional as opposed to bidirectional stimulation, and an asymmetric electrode geometry with the anode considerably larger than the cathode, thereby concentrating current density near the exiting cathode. If a similar stimulation paradigm is used, focal electrically-administered seizure therapy (FEAST) and focal electrically-administered therapy (FEAT) may be practical convulsive and nonconvulsive forms of rTES.
There has been no systematic evaluation of the components of electrical stimulation that may impact on the intensity, shape and focality of intracerebral charge density. There is a need to develop an understanding of how variation in the parameters of electrical stimulation (intensity of stimulation, electrode geometry, size, positioning, direction of current flow, etc.) impact on patterns of stimulation (e.g., peak intensity, degree of focality, perceived directionality, etc.). In other words, we need to develop better understanding of the basic principles that shape the patterns of stimulation. The present study investigates how various parameters in electrical stimulation might be manipulated in order to characterize general principles on how variation in electrical parameters impact on outputs. The purpose of this study is to develop a set of rules, that might apply to stimulation in more complex biophysical environments. Specifically, we examined how variation in the perceived quality and location of pain is a function of key parameters of electrical stimulation. This was studied under conditions intended to maximinze stimlation though the scalp (closer interelectrode distance), minimizing the contributions of Skull, CSF, and brain. It is hoped that the principles suggested here will provide a starting point for outlining the impact of these parameters on intracerebral current density.
Based on the literature cited above, we hypothesized that a unitary site of perceived pain would more commonly result from unidirectional than bidirectional stimulation as we believe that pain is experienced primarily at the point at which the electrical current exits the participant's head (i.e., one exit point with unilateral current, and essentially two exit points with bilateral current). Pain intensity would vary with the charge density at the cathode (i.e., a smaller cathode would be associated higher charge density and thus more pain). It was also hypothesized that the perceived location of pain would be largely determined by the location of the cathode (i.e., the primary location of the exit point of the electrical current).