Immediate neurophysiological effects
of transcranial electrical stimulation
Anli Liu 1,2, Mihály Vöröslakos3,4, Greg Kronberg5, Simon Henin 1,2,
Matthew R. Krause 6, Yu Huang5, Alexander Opitz7, Ashesh Mehta8,9,
Christopher C. Pack6, Bart Krekelberg10, Antal Berényi 3, Lucas C. Parra 5,
Lucia Melloni1,2,11, Orrin Devinsky1,2 & György Buzsáki 4
Noninvasive brain stimulation techniques are used in experimental and clinical ﬁelds for their
potential effects on brain network dynamics and behavior. Transcranial electrical stimulation
(TES), including transcranial direct current stimulation (tDCS) and transcranial alternating
current stimulation (tACS), has gained popularity because of its convenience and potential as
a chronic therapy. However, a mechanistic understanding of TES has lagged behind its
widespread adoption. Here, we review data and modelling on the immediate neurophysio-
logical effects of TES in vitro as well as in vivo in both humans and other animals. While it
remains unclear how typical TES protocols affect neural activity, we propose that validated
models of current ﬂow should inform study design and artifacts should be carefully excluded
during signal recording and analysis. Potential indirect effects of TES (e.g., peripheral sti-
mulation) should be investigated in more detail and further explored in experimental designs.
We also consider how novel technologies may stimulate the next generation of TES
experiments and devices, thus enhancing validity, speciﬁcity, and reproducibility.
Electrical stimulation to the brain has a long history in both science and medicine. In 1867,
Helmholtz discussed how electrical currents applied to the head can generate visual sen-
sations1. Currents as low as 0.3 mA could induce phosphenes, and stronger currents could
induce brighter and more lasting visual effects2. Subsequent studies conﬁrmed that these visual
phenomena result from retinal rather than direct brain stimulation3,4.
Early noninvasive electrical stimulation technologies used high intensities to directly affect
brain activity (see Box 1). Electroconvulsive therapy, introduced into psychiatry in the 1930s,
used currents of up to 60 mA to induce generalized seizures5. Subsequent studies on
DOI: 10.1038/s41467-018-07233-7 OPEN
1New York University Comprehensive Epilepsy Center, 223 34th Street, New York, NY 10016, USA. 2Department of Neurology, NYU Langone Health, 222
East 41st Street, 14th Floor, New York, NY 10016, USA. 3MTA-SZTE ‘Momentum’Oscillatory Neuronal Networks Research Group, Department of
Physiology, Faculty of Medicine, University of Szeged, 10 Dom sq., Szeged H-6720, Hungary. 4New York University Neuroscience Institute, 435 East 30th
Street, New York, NY 10016, USA. 5Department of Biomedical Engineering, City College of New York, 160 Convent Ave, New York, NY 10031, USA.
6Montreal Neurological Institute, McGill University, Montreal, QC H3A 2B4, Canada. 7Department of Biomedical Engineering of Minnesota, 312 Church
St. SE, Minneapolis, MN 55455, USA. 8Department of Neurosurgery, Hofstra Northwell School of Medicine, 611 Northern Blvd, Great Neck, NY 11021, USA.
9Feinstein Institute for Medical Research, Hofstra Northwell School of Medicine, 350 Community Drive, Manhasset, NY 11030, USA. 10 Center for Molecular
and Behavioral Neuroscience, Rutgers University, 197 University Avenue, Newark, NJ 07102, USA. 11 Max Planck Institute for Empirical Aesthetics,
Grüneburgweg 14, 60322 Frankfurt am Main, Germany. Correspondence and requests for materials should be addressed to A.L. (email: email@example.com)
or to G.Bák. (email: firstname.lastname@example.org)
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | www.nature.com/naturecommunications 1
electroanesthesia and electrosleep used subconvulsive current
intensities, delivered through large copper or platinum disks
covered with saline-soaked gauze over frontal and occipital scalp,
to affect large neocortical regions. To induce anesthesia, up
to 40 mA direct currents (DC) or alternating currents (AC; 1
Hz–200 Hz) were used (See Box 1), while weaker intensities
(3–10 mA) were typically needed for sleep induction6. More than
500 human surgeries were carried out under electrical anesthesia
supplemented with medication7, but serious side effects led to the
decline of the technique.
The remaining present-day application of high-intensity tran-
scranial electrical stimulation (TES, see Box 1) is for intra-
operative neuromonitoring. This technique, introduced by Mer-
ton and Merton (1980), uses high-intensity stimulation (up to
2000 V) through a pair of electrodes positioned over primary
motor cortex to generate a visible twitch in the contralateral hand
to monitor the functional integrity of central motor pathways
during resective surgery8–10. These early applications for high-
intensity TES6,11,12 laid the groundwork for subsequent experi-
ments which suggested that weak currents applied to the scalp
can also induce behavioral effects but without side effects and
without conscious awareness of the stimulation13,14. Compared to
other noninvasive neuromodulatory techniques such as tran-
scranial magnetic stimulation (TMS) and ultrasound stimula-
tion15, advantages of TES include low cost, portability, and
potential in-home applications, fueling a proliferation of human
trials16,17. However, a disadvantage of TES is that it may activate
excitable peripheral elements between the scalp electrodes,
including trigeminal nerve branches, the greater occipital nerve,
retina, and vestibular organs.
Despite more than 4000 publications (PubMed) on TES in the
past decade, we lack a mechanistic understanding of the
mechanism (or mechanisms) by which this technique produces
beneﬁcial or deleterious effects. Most TES studies place an elec-
trode above a targeted cortical region with the assumption that
the underlying neuronal activity will be boosted or suppressed.
While experiments in intact animals and computational models
have studied physiological and behavioral effects of TES, the
mechanisms are incompletely understood and their translation to
humans even more uncertain. It is also unclear whether TES
works through direct or indirect effects (see Box 1). This
mechanistic uncertainty explains why behavioral effects in
humans are often weak, variable and difﬁcult to replicate18,19.
In this Review, we discuss the immediate (i.e., acute, con-
current, or online, see Box 1) neurophysiological effects of elec-
trical ﬁeld stimulation in brain slices, rodents, non-human
primates, and humans. We compare experimental ﬁndings with
computational models, highlight the gaps in understanding, and
discuss novel methods to deliver spatially precise stimulation at
higher intensities than conventionally used, while reducing
unwanted peripheral effects. We close with recommendations for
future clinical and translational TES investigations. We do not
discuss here potential ofﬂine effects of TES, i.e. those effects
which occur or persist after the current has been switched off (see
Box 1). These have been reviewed elsewhere20,21.
Biophysics of induced electric ﬁelds
At each point within the brain, a scalar electric potential can be
measured relative to an arbitrary reference and expressed in Volts
(V). The electric ﬁeld
Eis the local change (gradient) of the
Eis a vector whose amplitude is measured in Volts per
meter (V/m). All transmembrane currents from nearby neuronal
processes contribute to generate E
When a current is applied to brain tissue (as in TES), it affects
the polarization of cellular membranes, which in turn can alter
neuronal excitability. Terzuolo and Bullock23 ﬁrst demonstrated
this using DC stimulation applied with metal bars to the crayﬁsh
abdominal receptor and lobster cardiac ganglion neurons.
Induced electric ﬁelds as low as 1 V/m affected the timing of
action potentials while much higher ﬁelds triggered action
potentials in silent neurons23. However, the orientation of the
ﬁeld was also important: rotating the applied electric ﬁeld’s
orientation relative to the main neuron’s soma-dendritic axis
(e.g., from parallel to perpendicular to the main neuronal axis)
affected the strength and direction of the effects. These ﬁndings
Box 1 Terminology
Transcranial electrical stimulation (TES)—The technique of applying an electric ﬁeld at the scalp surface with the purpose of directly affecting brain
activity. Early efforts applied ﬁelds at high intensities for electroconvulsive therapy (ECT, 60 mA), electroanesthesia (40 mA), and electrosleep (3–10 mA).
More recent efforts have applied TES at lower intensities (typically 1-2 mA) to reduce peripheral side effects, such as skin sensation and phosphenes.
Direct current stimulation—Electrical current applied in a constant, unidirectional manner ﬂowing from anode to cathode. When DC stimulation is
applied across the scalp, it is commonly referred to as transcranial direct current stimulation (TDCS).
Alternating current stimulation—Electrical current applied in a varying, typically sinusoidal, waveform, with current ﬂowing from anode to cathode in
one half-cycle and in the reverse direction in the second half-cycle. When the AC stimulation is applied across the scalp, it is commonly referred to as
transcranial alternating current stimulation (TACS).
Electric ﬁeld—The difference in voltage between two locations (in the brain). The electric ﬁeld is a vector with both magnitude and direction, and is
measured in units of Volts/meter (V/m).
Direct effects—Potential effects of TES on the excitability of neurons in the brain, which include both immediate and cumulative effects.
Indirect effects—Potential side effects of TES on non-neuronal elements, including placebo, peripheral nerves, retina, cochlea, glia, immune system, and
blood ﬂow as well as cumulative, long-term metabolic effects on neurons.
Immediate effects—Potential effects of TES which occur simultaneously or acutely during stimulation. These effects are measured by changes in
neuronal ﬁring patterns or the local ﬁeld potential (deﬁned below). We use the terms acute, concurrent, or online as synonyms of the term “immediate.”
Ofﬂine effects—Potential effects of TES which may outlast the period of stimulation.
Local ﬁeld potential—a composite measure of electric current generated from all active cellular processes within a volume of brain tissue superimpose
at a given location in the extracellular medium and generate an extracellular potential V
(a scalar measured in Volts) with respect to a reference
potential. The difference in V
between two locations gives rise to an electric ﬁeld (a vector, V/m). The local ﬁeld is monitored by small intracranial
electrodes as opposed to those obtained by scalp EEG recordings.
Oscillations—Rhythmic ﬂuctuations in the local ﬁeld potential, which can range in frequency from ultraslow (0.05 Hz) to ultrafast (500 Hz), with
speciﬁc oscillation frequencies characterizing particular brain states.
Shunting—During TES application, the ﬂow of current away from the brain due to passage along a low-resistance pathway (skin, subcutaneous tissue),
and away from a high resistance pathway (the skull). Current shunting explains why only a small part of the current delivered through TES electrodes
actually reaches the brain surface.
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generalize to the mammalian brain. An anode on the pial surface
(cathode at depth) depolarizes neurons and increases neuronal
ﬁring frequency, while the reversed current ﬂow (surface cathode,
depth anode) hyperpolarizes neurons and reduces their ﬁring
Yet, the above picture is an oversimpliﬁcation. First, it assumes
that the soma is singularly affected, but axon initial segment,
dendrites, and axon are also affected with different, unknown
sensitivities to electric ﬁelds. These combined and variable effects
on neuronal compartments can alter neuronal spiking19.In
addition, tonic depolarization of axon terminals could release
neurotransmitters in an axon branch-speciﬁc manner without
generating action potentials. Second, although ﬁelds perpendi-
cular to the soma-dendritic axis may have little inﬂuence on the
apical dendrites of pyramidal neurons26–28, they can activate
basal dendrites and dense local axonal arbors of basket and
chandelier interneurons. In addition, neurons with star-shaped
dendrites—many inhibitory interneurons, layer 4 stellate cells,
thalamocortical cells, and basal ganglia neurons—might be
affected equally by all ﬁeld orientations. These complexities
(Fig. 1) are relevant because the orientation of TES-induced
intracranial ﬁelds relative to these different neuronal populations
and compartments varies considerably29. In summary, the
inﬂuence of TES current depends not only on the amplitude but
upon how the induced electric ﬁeld relates to the three-
dimensional organization of neuronal compartments28,30,31.
Biophysical experiments and detailed computational models are
needed to fully understand these complexities.
Hypothesized immediate effects of electric ﬁelds on neural
We deﬁne immediate effects of TES as those that occur simul-
taneously or acutely from the modulation of the membrane
potential by the external ﬁeld. For a sufﬁciently strong electric
ﬁeld, neuronal ﬁring rates should result from changes in a neu-
ron’s input/output function19, facilitated neurotransmitter release
from presynaptic terminals32, or ectopically induced spikes33.
With AC stimulation, current ﬂows from anode to cathode in
one half-cycle and reverses direction during the other half-cycle.
These half-cycles ﬁelds are equal magnitude but opposite direction.
Thus, during low-frequency AC stimulation, somata may be
alternately depolarized and hyperpolarized, which could mimic
effects of anodal and cathodal DC stimulation. If the applied AC
frequency is much faster than the neuronal membrane time
constant (~30 ms, or > 33.3 Hz), the fast-changing polarities of the
stimulation may reduce neuronal polarization34,35, and thereby
reduce unwanted peripheral nerve stimulation effects (see below).
In tACS studies, applied TES current is posited to affect the
brain’s native oscillatory patterns36–38 and thereby target cogni-
tive and therapeutic applications. While the magnitude of TES-
induced physiological effects varies with the strength of the
induced ﬁeld, the ‘effectiveness’of the stimulation and its con-
sequences on neuronal and network activity depend on many
other experimental variables and the brain’s state. We distinguish
ﬁve neural mechanisms that could affect network activity: sto-
chastic resonance, rhythm resonance, temporal biasing of neu-
ronal spikes, entrainment of network patterns, and imposed
patterns (Fig. 2). These ﬁve mechanisms of ﬁeld-induced effects
can cooperate or compete with each other and with endogenous
brain activity, and can occur simultaneously in different networks
of the same brain.
(1) Stochastic resonance. Because physiological (EPSP, IPSP)
and exogenous (TES) polarizing mechanisms are added, there is
no theoretical ‘minimum effective threshold’of the induced
electric ﬁeld39,40. When a neuron nears the threshold of spike
generation, a very small amount of applied ﬁeld can bias spike
timing or spike probability. This is known as ‘stochastic reso-
nance’41. Weak stochastic effects are hard to quantify in normal
networks with ﬂuctuating activity levels because the affected
neurons may be scattered in distant networks. Upon repeated
application of the ﬁeld, different neurons may be excited or
hyperpolarized, and thus it may be difﬁcult to generate repro-
ducible effects based on this mechanism alone. (2) Rhythm
resonance. Under well-controlled, closed-loop conditions, a very
weak ﬁeld can be precisely timed to the depolarizing phase of an
oscillating neuron’s membrane potential (see below). Even
without a closed-loop system, rhythm resonance may occur when
an AC ﬁeld is applied to a regular endogenous rhythm at
the same frequency, because the external ﬁeld can affect the
native oscillation at a similar phase during each cycle42–50. Such
phase matching may explain tACS beta oscillation phase-
dependent modulation of TMS-induced motor responses in the
motor cortex51. (3) Temporal biasing of spikes. Strong rhythmic
applied ﬁelds may reliably affect the membrane potential and,
consequently, spike timing of subsets of neurons. This mechan-
ism is related to stochastic resonance since the coincidence of
intrinsic and extrinsic polarization forces work together but the
same neurons are more reliably activated from trial to trial. Such
phase-locked temporal biasing of neuronal spikes by the rhythmic
+ + + + + + + + + + + + + + + + +
– – – – – – – – – – – – – – – – –
Fig. 1 The impact of orientation of neuronal compartments on TES-induced
excitability. Four idealized neurons are shown with different orientation
relative to the induced electric ﬁeld. The electric ﬁeld can affect the soma,
dendrites, axon initial segment and the axon tree differently. The
relationship between the electric ﬁeld vector and the morphology/
orientation of neurons and individual neuronal compartments determine
whether the neuron will be net depolarized or hyperpolarized
Field intensity (V/m)
Fig. 2 Five postulated mechanisms to affect online spiking of neurons and
networks patterns in response to different estimated magnitudes of TES.
While the ﬁgure illustrates distinct effects, in reality the boundaries of
mechanisms are blurred under most experimental conditions. Several
mechanisms can act simultaneously in different networks of the same
brain. The numbers on the vertical axis are merely estimates based on
current data and theoretical considerations
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NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | www.nature.com/naturecommunications 3
forced ﬁeld may occur without affecting the overall ﬁring rates of
neurons. (4) Network entrainment. To entrain network patterns
of less regularity (e.g., slow oscillations of sleep) requires stronger
currents since exogenous patterns compete with native brain
rhythms19,40,45. (5) Imposed Pattern. Imposing an arbitrary
pattern on a neuronal network (e.g., enforcing theta activity on a
network with an endogenous beta rhythm) requires the strongest
ﬁeld to overcome the endogenous control of network neurons
Spike and local ﬁeld potential (LFP) measurements are often-
used methods of measuring neural activity in vivo, each with
advantages and disadvantages. Spike measures are the most direct
read-out of TES effect on neural activity35, but are not always
available in animals and are rarely possible in humans. LFP is the
extracellular measure of transmembrane electric current gener-
ated synchronously by a nearby population of neurons, brought
about primarily by postsynaptic potentials52,53 (see Box 1).
Importantly, LFP measurements allow comparison between
human and animal experiments. The acute effects of tACS could
be measured by changes in the LFP amplitude, such as epilepti-
form spike-and-waves or the dominant alpha rhythm as a func-
tion of the tACS phase35,54. A related method is to measure
entrainment of slower rhythms by tACS (e.g., sleep slow oscilla-
tions, see Box 1), also measured by the amplitude of faster, cross-
frequency coupled patterns (e.g., sleep spindles)17,25 Quantifying
entrainment of rhythms at the same frequency of the applied TES
is difﬁcult with current methods due to the difﬁculty of separating
the artifact from the entrained physiological pattern, as discussed
in more detail below.
Acute effects of electric ﬁelds in vitro
In brain slices, the effects of induced ﬁelds can be studied with
exquisite control, since all synaptic signaling can be blocked
pharmacologically55. Most studies apply constant ﬁelds across
parallel wires outside the slice. This spatially diffuse stimulation
differentially affects the neuronal compartments (e.g., soma,
apical dendrites), producing complex effects23. In vitro recordings
of rat hippocampal and neocortical pyramidal cells show that DC
electric ﬁelds of 1 V/m can polarize the soma of pyramidal neu-
rons by ~0.2 mV.
Yet, extracellular electric ﬁelds abound in the brain22. Local
electric ﬁelds (> 0.5 V/m) applied in vitro with a micropipette
near layer 5 pyramidal cell dendrites and soma induce small
intracellular voltage ﬂuctuations (e.g., 0.2 mV) and may phase-
lock the spikes to the low-frequency (1–30 Hz) oscillating external
ﬁeld without affecting overall spiking rate56. Thus, under idea-
lized conditions, very weak but localized extracellular ﬁelds can
bias neuronal activity (Fig. 3). Spike timing is highly sensitive to
small, persistent oscillations that act throughout a volume. This
voltage ﬂuctuation is comparable to intrinsic, non-synaptic,
internally generated noise in layer 5 pyramidal neurons (0.2–0.4
mV)57 and is much smaller than synaptic background activity
in vivo33,58. For example, hippocampal theta oscillations, for
instance, induce ~4 V/m electric ﬁelds across the CA1 pyramidal
layer, while hippocampal sharp waves induce 5–15 V/m39.
Endogenous ﬁeld strengths associated with slow oscillations in
the neocortex are 1–2 V/m45. Whether weak applied electric ﬁelds
can overcome such strong endogenous ﬁelds remains an impor-
tant experimental question.
AC ﬁelds of sufﬁcient strength applied at low frequencies
(< 10 Hz) can modulate ﬁring rate periodically by the same
mechanisms as DC ﬁelds59. For higher AC stimulation fre-
quencies, the induced effects are diminished; for example, somatic
polarization is reduced by ~70% when a 50 Hz AC stimulation
induces a 1 V/m ﬁeld, because the membrane cannot follow the
rapid ﬂuctuations in the electric ﬁeld60. The oscillatory ﬁelds may
0 0.2 0.4 0.6
0 2.1 4.2
V amplitude (mV)
field strength (mV/mm)
V amplitude (mV)
field strength (mV/mm)
0 0.2 0.4 0.6
0 2.1 4.2 6.4
Fig. 3 Field-entrainment of spikes under idealized conditions in vitro. Synaptic transmission was blocked pharmacologically. aFour neurons with somata
located within 100 μm of tissue were patched with intracellular electrodes (blue). Seven extracellular electrodes monitored extracellular voltage (Ve)
ﬂuctuations (magenta). An extracellular stimulation electrode (S1) was placed 50–80 μm from the recorded somata. bEach of the four intracellularly
recorded neurons were depolarized to induced spiking. Spiking in the absence (top traces) and in the presence of extracellular stimulation (magenta;
100 nA at 1 Hz; bottom traces) is shown. cSpike ﬁeld coherence (circles, mean; error bars, s.e.m.) between spikes and extracellular Veis shown during
1 Hz extracellular stimulation (black) and control condition (cyan; circles indicate mean Ve amplitude at the soma and error bars indicate s.e.m.) as a
function of stimulation strength. Asterisks indicate statistical signiﬁcance of the spike-ﬁeld difference between control and extracellular stimulation.
dSpike-spike coordination (spike-ﬁeld coherence for two simultaneously occurring spikes; essentially spike synchrony among neurons) during 1
Hz extracellular stimulation (black) and control (cyan) condition. Note that stronger ﬁelds are needed for coherent entrainment of neuronal spikes across
the four neurons (d) than for inducing spike-ﬁeld coherence for each neuron separately (c). Figure reproduced with permission56
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not affect the overall ﬁring rate but can modulate the timing of
ongoing spikes56,59 (Fig. 3). AC stimulation at a frequency
matching the resonant neuronal properties can induce cumulative
effects over multiple cycles43,45,59–62. Under such ideal condi-
tions, AC ﬁelds as low as 0.2–0.5 V/m can shift spike
timing23,43,59,60. Higher stimulation frequencies require stronger
currents to bias spike-LFP coupling (> 0.7 V/m induced ﬁeld at 1
Hz and > 5 V/m at 30 Hz56,39, Fig. 3), due to the frequency ﬁl-
tering properties of the neuronal membrane63. One caveat of
these in vitro studies is that pharmacologically or intracellularly
induced oscillations are much more regular than in vivo oscilla-
tions19. Overall, in vitro studies have laid the biophysical foun-
dation for TES, providing quantitative relationships between the
effective magnitude of voltage gradients for both V
activity. These experiments have also clariﬁed the geometric
relationship between the vectorial ﬁeld and neuronal morphology
and soma-dendritic orientation. However, these studies do not
reveal how applied current affects neurons when applied through
the skin, subcutaneous tissue, skull, dura and cerebrospinal ﬂuid
in the living animal. For a detailed summary of in vitro experi-
ments measuring acute physiological changes during TES, see
Supplementary Table 1.
Acute effects of external electric ﬁelds in rodents
In rodent experiments, electric ﬁelds are often applied with
electrodes placed on the skull, the dura mater, or directly on the
brain surface. Often, a saline-ﬁlled cup is applied at various epi-
cranial locations with a return electrode on the thorax or the
contralateral hemisphere. Several in vivo studies have found a
dose-dependent relationship between the induced electric ﬁeld
and spiking rate24,34,35,40, transmembrane potential and, at
higher stimulation intensities, on the LFP25,35,39. Applied electric
ﬁelds can alter neuronal excitability64 and evoked responses64,65
in a polarity-speciﬁc manner.
However, rodent studies typically use ten-fold stronger current
intensities compared to human studies66. Across the 28 rodent
experiments reviewed67 (Supplementary Table 2), the induced
intracranial ﬁelds averaged 6.8 ± 3.8 V/m (n=11, ten epicranial
and 1 subdural studies), compared to < 1 V/m measured in con-
ventional human TES studies19,68,69. While stimulating transcra-
nially in rats, the lowest electric ﬁeld sufﬁcient to affect the timing
of spiking activity in widespread cortical and hippocampal areas
was ~1 V/m; higher intensities were required to reliably affect LFP
and the membrane potential in intracellularly recorded neurons
in vivo35,40.Higherﬁeld intensities were also required in urethane-
anesthetized rats to affect LFP oscillations70,71. Similarly, applied
currents have terminated thalamocortical spike and wave patterns
in rodents (1.5–2 mA, applied to skull)54,72 or triggered paw
movement (0.25 mA and 0.9 mA)34 and modulated endogenous
slow wave initiation and propagation patterns (~3 mA /mm2)73.In
summary, rodent stimulation studies demonstrate physiological
effects on spike timing, LFP oscillations, and terminating seizure
patterns, but applied currents were several-fold stronger than the
currents expected to penetrate the human brain using TES. For a
detailed summary of in vivo experiments measuring acute phy-
siological changes during TES, see Supplementary Table 2.
Immediate electrophysiological effects of TES in primates
Non-human primates permit TES investigations using a relatively
large head with a gyrated cortical surface. The variable anatomy,
including thick skull and large musculature, can result in a wide
range of ﬁeld strengths across studies and even across individuals.
Opitz et al.69 measured ﬁeld strengths in two anesthetized cebus
monkeys with small round scalp electrodes (3.14 cm2) over the
left occipital and frontal cortices stimulated with 1–2 mA. The
median induced electric ﬁelds were 0.21 V/m (strongest 0.36 V/
m) in a male monkey (4.1 kg) and 0.39 V/m (strongest 1.16 V/m)
in a female monkey (2.9 kg). The weaker effects in the male likely
reﬂected the larger head musculature that shunted much of the
Krause et al.74 studied the effects of tDCS with intracranial
electrodes in right prefrontal and left inferotemporal cortices in
macaque monkeys performing an oculomotor foraging task. Based
on a ﬁnite-element model75, 2 mA anodal stimulation applied to the
right prefrontal region induced ﬁeld strengths between 0.4–0.7 V/m.
Although tDCS did not affect the spontaneous or task-related ﬁring
rates of isolated single units or multi-unit clusters, LFP power in the
1–15 Hz band increased ten-fold during DC stimulation, and
coherence within the prefrontal array increased at all frequencies
from 2–100 Hz. However, the changes in spike frequency or spike-
ﬁeld coherence in areas with stronger ﬁelds (up to 1 V/m) but
outside the recording area, could not be ruled out.
Kar et al.76 delivered tACS of 1 mA peak amplitude at 10 Hz
frequency through scalp electrodes placed anterior to the vertex
and lateral to the middle temporal area (MT). In one monkey, the
induced ﬁeld was 0.12 V/m at closely spaced electrodes (0.5 mm)
in the recorded MT area. Spiking activity during the stimulation
was not reported, but a reduction in spike adaptation to the visual
input after tACS offset was observed.
In a dose-ranging tDCS experiment on motor control, Lee et al77.
applied anodal stimulation over primary motor cortex (with cathode
at the vertex) to determine the threshold to elicit a muscle twitch in
sedated macaques. Between 50–120 mA (~35–55 V/m electric ﬁelds
in the motor cortex) were needed to trigger a visible twitch.
In summary, primate models offer the advantage of testing TES
in a larger head model with a gyrated cortical surface. The few
primate studies which have been performed to date employ dif-
ferent montages and measurement techniques. Compared to the
rodent models, primate experiments are limited by weaker
induced ﬁelds and proportionately less cortical recording cover-
age. Primates with large heads, thick skulls, and greater muscu-
lature have greater shunting, or current loss away from the brain,
and require larger current intensities to achieve comparable
electric ﬁeld strengths to those in rodent brains. These differences
highlight the critical need to match intracerebral ﬁelds rather
than stimulation parameters across species and experiments.
Ex vivo human studies
Translating the results of in vitro and in vivo animal experiments
to humans is confounded by complexities of brain and head
anatomy. Neurons in highly gyrated cortices of higher primates
have variable alignments even on the convexity29,78, in contrast to
those of lissencephalic rodents which are mostly perpendicular to
the skull surface. Since entrainment efﬁcacy of neurons depends
on how polarized neurons align with the direction of the electrical
ﬁeld, TES of human cortical gyri may lead to unpredictable
effects. Further, safety considerations require TES application
through the hair, skin, and skull. In contrast, few animal
experiments applied TES through the scalp35,69,74,76. Using
transcranially applied stimulation, only a small fraction of the
applied current enters the brain: most is lost through the skin,
subcutaneous soft tissue, and the skull35 (Fig. 4a). One study
conducted in rodents and human cadavers demonstrated that >
75% of applied TES current is lost through shunting35, although
the shunting effects differ between humans, nonhuman primates
and rodents in unknown ways.
Human cadavers yield insights regarding conductivity and
anisotropy. Multiple intraparenchymal, three-dimensional mea-
surements of the electric ﬁeld distribution can be obtained and
multiple stimulation parameters can be tested in the same cadaver
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(Fig. 4b, c). Such empirical measurements help determine the
basic biophysical tissue properties and help model effects of TES
into the brain parenchyma79–81 (Fig. 4d–f). Intracerebral electric
ﬁelds were measured in intact cadaver brains at 198 sites in a
three-dimensional grid, while applying stimuli that varied in
frequency, intensity, phase relationship, electrode location, and
size35. The authors concluded that scalp, skull, brain, and cere-
brospinal ﬂuid behave as ohmic conductors. In other words,
induced ﬁelds increased linearly with stimulation intensity, but
were independent of stimulation frequency82. The comparison of
transcutaneous, subcutaneous and epidural stimulations of the
cadaver brain led to the conclusion that the bilateral transcuta-
neous stimulation of the cadaver heads with traditional montages
and intensity (peak 1 mA) induced only ~0.2 V/m ﬁeld in the
brain35. Of note, death imparts profound changes in the bio-
physical properties of brain tissue, limiting a direct ex vivo to
in vivo comparison83.
In vivo human studies
A critical question is the magnitude of the electric ﬁeld reaching
the brain when using conventional TES, applied transcutaneously
at peak intensity currents of 1–2 mA, because ﬁeld strength
constrains the expected nature of the physiological changes
(Fig. 2). The main barrier for TES to reach the brain is the skull’s
high resistance (~160 Ωm) combined with the low resistance of
the scalp (~2 Ωm)80. A large fraction (up to 75%) of the current
seeks the path of least resistance and shunts across the scalp35.
Three independent studies using intracranial electrodes in epi-
lepsy patients undergoing surgical evaluation demonstrated that
maximal electric ﬁelds at the cortical surface under the electrodes
were < 0.5 V/m for 1 mA peak intensity currents68,69,84.
Can such weak electric ﬁelds reaching the human brain induce
changes in neuronal networks, as measured by changes of the
LFP? Oscillations are ubiquitous in the human brain, ranging
from ultraslow (0.05 Hz) to ultrafast (500 Hz), with speciﬁc
oscillation frequencies characterizing different arousal states or
cognitive processes and coordinating activity between local and
distant brain regions85. A leading hypothesis about the behavioral
effect of tACS is that the applied stimulation frequency matching
the dominant network oscillation entrains and augments endo-
genous rhythms37,38,62,86–89, such as gamma60,89,90, beta91,
alpha46, theta92, and slow frequency oscillations40,45,93.
Whether the weak electric ﬁelds resulting from TES at con-
ventional intensities (1–2 mA peak) can entrain neuronal oscil-
lations is technically difﬁcult to assess. One technical challenge is
that the magnitude of the stimulation artifact is dramatically
larger than the magnitude of the endogenous oscillation, making
artifact rejection difﬁcult. To illustrate, supposing the use of a
10 Hz tACS waveform to entrain endogenous alpha oscillations,
the amplitude of alpha waves is ~ 100 µV over the parietal-
occipital lobes, whereas the artifact induced by a 2 mA tACS
stimulus can exceed 1 V (~1000x greater). Therefore, even if
99.9% of the artifact is removed, the remaining 0.1% of the
artifact would be as large as the endogenous alpha rhythm.
Further, the stimulus waveform can be distorted, resulting in
multiple harmonics of the applied signal. Due to non-linearities
and instability of the stimulation waveform during the experi-
ment, the waveform of the stimulation artifact and its harmonics
across trials cannot simply be subtracted from the resulting EEG
recordings. While artifact rejection algorithms can clean the
recorded signals, artifact-free brain activity is unlikely to be
recovered when the stimulation frequency matches the endo-
genous oscillation frequency94. Another challenge is the
Electric field (mV/mm)
Electric field (mV/mm)
Electric field (mV/mm)
Relative shunting effect
electrode Skin Skull Brain
Distance from skin surface (mm)
00 0.5 1 1.5 6
10 100 1000
Stimulus frequency (Hz) Stimulus intensity (mA)
10 50 100
Stimulus intensity (µA)
6 × 10 mm
5 × 15 mm
Fig. 4 Current loss of TES in rodents and in human cadaver brains. aCompared to subcutaneous stimulation (red), transcutaneous stimulation (blue)
generated several-fold weaker electric ﬁelds in a rodent model. bSchematic of the experimental arrangement for transcutaneous, subcutaneous, and
epidural stimulation in cadavers, in a coronal plane. cPhotograph of the cadaver skull with drilled holes and inserted matrix of recording electrodes.
Stimulation electrodes, marked by blue and red circles for negative and positive polarity, respectively, were ﬁxed to the skull. dEffect of stimulus frequency
on intracerebral voltage gradients. Stimulus frequency between 5 and 1000 Hz has a minor effect on intracerebral gradients. eExtrapolation from
measurements in human cadavers (blue crosses) suggests that approximately 6 mA current applied across the skin would induce 1 V/m intracerebral
electric ﬁeld (blue open circle). fAttenuation of charge ﬂow (red line) through scalp (pink), skull (yellow), and brain (dark pink), as measured in human
cadaver heads. Figure adapted from Figs. 1,4,5in Voroslakos et al., (2018)35, under the Creative Commons License
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saturation of recording ampliﬁers, which precludes analysis of the
endogenous LFP during stimulation.
Besides neurophysiology, other experimental measures of TES
immediate effects include blinded subject reporting, TMS-evoked
motor evoked potential (MEP), and fMRI. Feurra et al.49 applied
at 1.5 mA anodal current at varying frequencies of (2–70 Hz) over
somatosensory cortex and asked subjects to report tactile sensa-
tion in the contralateral hand. While no sham or control sti-
mulation site was included, during one of the two replications,
subjects reliably reported sensation during alpha (10–14 Hz) and
high gamma (52–70 Hz) frequency stimulation. These occasional
behavioral effects of TES may reﬂect stochastic resonance (Fig. 3).
A study of TMS-evoked MEPs during tACS (1 mA peak-to-
peak) at different frequencies (theta, alpha, beta, and gamma) over
the primary motor cortex during rest and during a motor imagery
task found MEPs were greatest during beta frequency tACS at rest,
but during theta–tACS during motor imagery38. Guerra et al.51,
found beta frequency (1 mA peak-to-peak, 20 Hz) tACS stimula-
tion acutely modulated the MEP amplitude in a phase-dependent
manner; however another experiment37 (1.5 mA peak-to-peak)
did not ﬁnd phase-dependent modulation of corticospinal
responses using low-frequency tACS (0.8 Hz) or tDCS super-
imposed with 0.8 Hz. An fMRI study applied 20 Hz tACS stimu-
lation (1 mA peak-to-peak) over the primary motor cortex (M1)
found enhanced local connectivity within M1, without changing
the overall local activity or long-range connectivity within the
default mode network95. While these TMS and fMRI ﬁndings are
consistent with a rhythm-resonance mechanism (i.e., 20 Hz
modulation of resting beta frequency in motor cortex; Fig. 2), a
peripheral contribution remains possible. Both consciously
detected and subthreshold (unconscious) sensory stimuli can
inﬂuence neocortical activity96,97.
Epilepsy patients undergoing invasive monitoring offer the
opportunity to measure and analyze endogenous oscillations
during TES and thus directly monitor brain responses, since
intracranial signals are an order of magnitude larger than those
recorded from the scalp. The effects of tACS (0.75 Hz and 1 Hz)
during NREM sleep and waking rest were assessed in epilepsy
surgery patients at standard stimulation intensities (up to 2.5 mA
peak to peak; maximum induced ﬁeld: 0.43 V/m). During NREM
sleep, slow oscillations (~1 Hz) strongly entrain spindle activity in
the 10–15 Hz range. This provides the opportunity to test whe-
ther endogenous slow-wave rhythms can be entrained with low
frequency (0.75 Hz or 1 Hz) tACS, while measuring entrainment
in a different frequency band that is largely uncontaminated by
stimulation artifacts. While spindle and gamma activity robustly
entrained to the phase of the endogenous slow oscillation, low-
frequency tACS failed to entrain spindle or gamma activity at
over 1000 electrode sites measured in patients during NREM
sleep. Likewise, no entrainment of gamma or theta activity was
observed during waking rest19,98.
Instead of entrainment, TES-phase/EEG-amplitude coupling
can also be examined with scalp recordings. This approach was
validated by concurrent unit recordings in rat experiments54.In
agreement with the intracerebral recordings19, no detectable
effect was obtained on the scalp EEG when TES was applied at <
2.5 mA in healthy subjects35. However, reliable tACS phase
modulation of the amplitude of alpha rhythms was present when
intensities exceeded 4.5–6 mA. The effect was speciﬁc for neu-
ronal stimulation of cortical neurons because alpha amplitude
enhancement over the left and right occipital cortex occurred in
phase with alternating the anodal stimulation. Furthermore, to
rule out peripheral contributions, control experiments applying
stimulation to the abdomen did not produce an effect on alpha
Novel TES methods
Current applied through the scalp stimulates electrically sensitive
elements from the scalp surface to the brain. These peripheral
side effects, including stimulation of cranial nerves, retina, and
vestibular system, limit the maximum tolerable dosing in humans
using traditional TES approaches. Currents exceeding 1–2mA
(depending on electrode type) can cause itching, burning sensa-
tion, pain at the skin under and around the electrodes99.
The next generation of TES technologies should make three
advancements: (1) delivery of stronger currents to the brain while
minimizing peripheral and indirect effects; (2) simultaneous sti-
mulation and recording of brain activity for quantitative mea-
surements of TES-induced effects; and (3) targeting of speciﬁc
rhythms through closed-loop stimulation of brain areas, includ-
ing deep brain structures.
Multi-electrode stimulation is one potential method to increase
current density and focality of stimulation delivered to the brain
surface. The strongest transmembrane polarization is expected to
build up where electric ﬁelds overlap inside the head, whereas the
intensity of local scalp stimulation is divided proportionally by
the number of stimulation electrode pairs. The approach is
similar to previous efforts which used multiple small electrodes to
optimally target speciﬁc brain areas100.
Two other approaches attempt to increase the brain to scalp
stimulation ratio. The ﬁrst is temporal interference stimulation.
First introduced for stimulating peripheral nerves and brain101,
temporal interference is now routinely used in physical ther-
apy102. When applied to the mouse brain, intersecting 2 kHz and
2.01 kHz sinusoidal waveforms yielded a 10 Hz amplitude mod-
ulation of the LFP34. Because the high-frequency stimulus exceeds
the time constant of the peripheral nerve membrane, the fast
alternating current may reduce skin sensations and avoid retinal
responses, while the interfering signals in the brain are expected
to be strong enough to modulate neuronal ﬁring. The second
approach, ‘intersectional short pulse‘(or ISP) stimulation, uses
multiple electrode pairs35. This time-multiplexing method
exploits the time integrating property of the neuronal membrane
(i.e., the membrane time constant of cortical neurons is ~30 ms),
by applying currents that switch every 60 µs between the electrode
pairs. The strongest transmembrane charge will build up where
successively induced electric ﬁelds overlap inside the head,
whereas the intensity of local scalp stimulation is divided pro-
portionally by the number of stimulation electrode pairs. Again,
the more electrode pairs are used, the greater the potential to
generate brain focal stimulation while minimizing stimulation of
the scalp to reduce skin sensations78,100. Initial studies using
temporal interference and ISP show increased spatial targeting in
animal models, but further analysis and reﬁnements of these
techniques are necessary.
The application of high-frequency pulses avoids the saturation
of recording ampliﬁers and allows for the simultaneous mea-
surement of the direct physiological effects of tES in humans35.
Ampliﬁer modiﬁcations, such as front-end subtraction of the
applied tES waveform can further reduce induced artifacts. Such
innovations are essential to directly quantify the physiological
responses of neuronal circuits.
Recommendations for future experiments
Although the lasting effects of stimulation are the main purpose
of most clinical trials103,104, our review focused on the immediate
physiological effects of TES on spikes and network dynamics. A
more accurate and sophisticated understanding of physiology,
grounded in empirical measurements, will improve design and
execution of future TES experiments, with more potent and
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consistent clinical beneﬁts. Below, we offer a list of tasks for
future TES research.
Exploration of peripheral and non-neuronal mechanisms of
TES. It is tacitly or explicitly assumed that behavioral and clinical
effects of TES are mediated by directly affecting neuronal activity
in the brain. However, behavioral and cognitive changes in
response to scalp electric stimulation may be mediated by other
effects as well. Electric ﬁelds affect the excitability of sub-
cutaneous nerves which signal to the brain. Even when the subject
does not report a conscious experience of TES, sensory stimula-
tion can still indirectly affect brain circuits97,105. Therefore,
control experiments that measure the conscious and subconscious
effects of sensory stimulation should be included in future studies.
This can be achieved by including an “active control,”such as
stimulation of the neck or abdominal areas35 or application of
TES to another region of the brain106. In addition, applying
topical anesthesia to the scalp under the electrodes and testing
varying intensities of skin stimulation could exclude or char-
acterize sensory contributions.
Long-lasting TES effects can be indirectly mediated by non-
neuronal mechanisms as well, including trophic factors107,
neurotransmitter metabolism51,108 glia109,110, neuroendocrine
system111,ﬁbroblasts112 lymphocytes, and other electrically
charged immune system components113. These potential non-
neuronal effects should be considered with low-intensity scalp
stimulation, which may be too weak to induce the necessary
electric ﬁeld strengths to instantaneously affect neuronal activity.
These indirect effects may play a role even when neuronal activity
is directly affected by stronger TES.
Targeting neural circuits. Growing evidence, summarized above,
suggests that TES applied at ± 1 mA peak intensity (2 mA peak to
peak) generates < 0.5 V/m electric ﬁelds in the human brain. This
is sufﬁcient to generate 0.1–0.2 mV changes in the membrane
potential of cells within the stimulated area. As these changes are
signiﬁcantly lower than the 20 mV depolarization required to
bring a neuron from its resting potential to spike threshold
in vitro, TES is unlikely to directly elicit changes in spiking
activity. Instead, weak induced ﬁelds may be more effective when
used to bias or augment ongoing rhythms instead of introducing
new activity patterns. Applying current at the optimal phase of
endogenous rhythms in a closed-loop system may be most
effective54. However, such responsive methods require character-
ization of the affected circuitry and continuous monitoring
and adjustment of the relevant rhythms114. This may be achieved
with weak electric ﬁelds but other applications may require
higher ﬁeld strengths. For example, prompt control of spiking
activity (e.g., to terminate a seizure) may require ﬁeld strengths >
Modeling for spatial targeting and to increase replication
across subjects and studies. Individual variations in neuroa-
natomy (scalp, muscle, and skull thickness) can signiﬁcantly alter
the strength and distribution of the induced electrical ﬁeld in
humans68,98,115–117 and experimental animals75,77, and thus may
account for inconsistent behavioral results in human TES
experiments. This is exempliﬁed by Fig. 5, which compares two
individual models, with the same surface electrode montage but
different head anatomy (Fig. 5a, b, created using previously
published data and methods118,119). Differences in skull thickness
and cerebrospinal ﬂuid volume in posterior regions lead to dif-
ferent ﬁeld distributions (Fig. 5a, b).
Computational models of current ﬂow have been continuously
reﬁned over the last decade78,100,116,117,120–122. Recent head
models can improve spatial targeting by guiding stimulation
electrode placement (Fig. 5c). These models are based on
segmentation of tissue (at 1 mm3resolution) into compartments
of differing electrical conductivities, including skin and other soft
tissues, skull, gray matter, white matter, ventricles, and cere-
brospinal ﬂuid. Models indicate that focal stimulation of a few
centimeters may be achieved at the cortical surface100, generally
with a trade-off between focality and intensity of stimulation
(Fig. 5). Deep targets may be reached at relatively high ﬁeld
intensities when adjacent to CSF-ﬁlled ventricles, which can guide
currents deep into the brain68.
Computational models can help optimize electrode locations
for maximal stimulation of a desired brain target100. Individual
MRI head anatomy combined with toolboxes such as ROAST,
and SimNIBS, and SCIrun can deﬁne ﬁeld distribution for this
purpose. When an individual’s head MRI is not available,
validated and calibrated universal head models can help
reproduce in vivo intracranial measurements118,119,123. Further-
more, multielectrode montages can be optimized to engage local
and network level targets, as demonstrated by a technique
combining electric ﬁeld modeling and cortical networks derived
from PET or resting state functional connectivity MRI124.
Human TES experiments typically report the stimulation
protocols by reporting the stimulating electrode positions (e.g.,
“anode at F8”or “over left dorsolateral PFC”) and current
intensities (e.g., “2mA”). These speciﬁcations describe how the
experiment was designed and performed, but they only indirectly
deﬁne the induced electrical ﬁeld strength and distribution.
Reporting induced ﬁeld estimates from head models would help
relate behavioral and clinical reports to underlying neural
mechanisms. When the individual MRI head is available,
subject-speciﬁc estimates of the induced ﬁeld strength should
be included as a covariate in analyses of its behavioral or neural
effects. Intensity-response curves, inverted polarity stimulations
and skin stimulations far from the desired target location may
distinguish between direct brain-stimulation and peripheral
effects. Measurements or estimates of the electric ﬁeld strength
are included in some recent human and large animal studies but
are missing from most rodent studies. Reporting ﬁeld magnitudes
is important for animal studies where variable head size and
anatomy can strongly affect ﬁeld magnitudes. These values would
provide a critical translational scaling metric to facilitate
comparisons with human and animal studies and should be
included in future publications.
Hardware and signal processing issues. The simultaneous
measurement of physiological activity (EEG, ECoG) during tACS
is complicated by stimulation artifacts several orders of magni-
tude larger than intrinsic brain activity (1–100 µV). When brain
activity and stimulation are recorded simultaneously, special care
should be taken to (1) avoid saturation of the recording ampliﬁer
during experiments, and (2) subtract stimulation artifacts from
the measured signal. Saturation can be avoided by using high-
quality ampliﬁers with sufﬁcient input dynamic range (e.g. an
ampliﬁer with dynamic range of ± 100 mV can capture the full
Artifact rejection methods initially suggested for scalp EEG87
are highly unlikely to fully recover artifact-free brain activity (see
above)94,125. Therefore, when using tACS stimulation, it is
difﬁcult to analyze online neural responses at the same frequency
in which stimulation is applied (e.g., increases in alpha
oscillations during 10 Hz tACS stimulation). Analysis should
target frequencies well away from the stimulation frequency and
its harmonics. Measurements using phase-amplitude coupling or
single-unit recordings may also be used when available.
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DC stimulation artifacts pose less of an issue during signal
processing as hardware ﬁlters will typically remove any DC offset
(avoiding ampliﬁer saturation). However, EEG recordings in
humans are affected by artifacts in the 1–15 Hz frequency band
due to heartbeats and eye-blink125,126. The effect is most
prominent in frontal locations. In addition, care should be taken
to avoid motion of the subject and stimulation electrodes,
electrical noise injected into the system is reduced by using high-
quality (i.e., narrow band) stimulation devices, and environ-
mental noise (ensure adequate electrical isolation19,68. Control
experiments, e.g., using a phantom head model, are needed to
separate TES-related artifacts from the induced neural activity.
Data sharing standards. Because different devices use different
conventions, it is critical to specify whether tACS amplitudes
describing the absolute magnitude (i.e., the amplitude of one
phase), the peak-to-peak amplitude, or another form of para-
meterization. Shared data and analysis methods can facilitate
coordinated community efforts and metanalysis. Combining
many data sets may offer novel insights and interpretations (e.g.,
https://www.nwb.org/). Further, given the artifact rejection
methods described above, sharing raw EEG data sets would
enhance replication of the signal processing pipeline. Sharing
analysis software packages will facilitate cross-validation of results
A more accurate and sophisticated understanding of TES phy-
siology, grounded in empirical measurements, will improve the
design and execution of neurostimulation experiments, yielding
more robust and consistent clinical beneﬁts. While the authors of
this review share a consensus on experimental recommendations,
we acknowledge a controversy within the TES ﬁeld and amongst
ourselves as to whether the attenuated electrical ﬁelds that reach
the human brain with conventional protocols can directly affect
neurons and neuronal networks. While direct neural mechanisms
have been measured with exquisite experimental control in vitro
and in vivo models, whether these insights translate to humans is
uncertain. We have highlighted challenges in translating
mechanistic insights from experimental models to humans,
including (1) the interaction of the applied external ﬁeld with
competing or cooperating network oscillations when comparing
in vitro to in vivo studies; (2) external ﬁelds applied to the folded
cortical surface produce more variable and unpredictable effects
in larger animals (including humans) compared to rodent and
in vitro models; (3) the large human head and the low resistance
scalp shunt most applied current, with much smaller fraction of
the applied current reaching the cortical surface.
The weak induced electric ﬁelds reaching the human brain
contrast with the numerous behavioral and clinical effects
reported. We should also consider how TES can affect brain
activity indirectly, including activation of afferent nerves127,
Fig. 5 Individualized models of transcranial electrical stimulation for two subjects with variable electrode placements. aModel of Subject 1, with 2 mA
current injected at electrode Fp1 and return from electrode P3. bModel of Subject 2 with the same electrode conﬁguration. cModel of Subject 2 with
another electrode conﬁguration attempting to achieve more focal stimulation. Note that stimulation is more focal but achieves weaker ﬁelds compared to
b, due to increased current shunting through the skin between near-by electrodes. Electrode montages in band care obtained by a numerical optimization
algorithm that attempts to achieve maximal intensity (b) or maximal focality (c) for the location indicated by a black circle. The current-ﬂow models were
generated from previously published data118 and methods119. We used ROAST, a toolbox for realistic current-ﬂow models of the human119
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retina and the vestibular apparatus42, astrocytes, perivascular
elements, and glial activation109,110, as well as through placebo
effects128. Both tACS and tDCS may also induce cumulative or
longer-term effects via unknown mechanisms that merit further
exploration. Future human behavioral trials should control for
sensory contributions with approaches such as topical anesthesia,
or by using sensory controls such as stimulation away from the
Novel stimulation methods that safely induce higher brain to
scalp electric ﬁeld ratios are needed. Existing experimental tech-
niques, such as temporal interference stimulation and intersec-
tional short pulse stimulation delivering stimulation at high
frequencies (> 1 kHz), require further testing in both experi-
mental animals and humans. These and related methods, such as
multielectrode TES78,100, may better target brain structures
compared to conventional electrode montages. Closed-loop
techniques may enhance the efﬁcacy of TES via resonant mod-
ulation of native brain rhythms. In selected cases, supported by
experimental ﬁndings and human head-based modeling, sub-
cutaneous arrays or electrode plates inserted below the skin may
provide superior efﬁcacy, similar to techniques to treat head-
ache129 and facial pain129,130 by stimulating peripheral nerves.
With further development, TES could become a leading tool for
chronic, on-demand, at-home treatment of many neuropsychia-
tric and neurological diseases.
Received: 26 April 2018 Accepted: 18 October 2018
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We thank NYU Finding a Cure Against Epilepsy and Seizures (FACES) for sponsoring
the Minisymposium Current state of Transcranial Electrical Stimulation: Progress &
Challenges, on December 18, 2017 where the discussions contained in this manuscript
arose. In addition, A.L., S.H., O.D., and G.B. receive support from R01 MH 107396.
A.L. and G.B. have composed and written the review with contributions from M.V., G.K.,
S.H., M.K., Y.H., A.O., A.M., C.P., B.K., A.B., L.P., L.M. and O.D.
Supplementary Information accompanies this paper at https://doi.org/10.1038/s41467-
Competing interests: A.B. is the founder and owner of Amplipex and Neunos LLCs,
which manufacture biosignal ampliﬁers and stimulator devices. A.B. and G.B. have ﬁled a
patent application about the ISP method. A.O. is an inventor on patents and patent
applications describing methods and devices for noninvasive brain stimulation. B.K. is a
paid consultant for Aqeel, LLC, which develops transcranial stimulation technology. L.P.
has shares in Soterix Medical Inc., which develops transcranial stimulation
technology. The remaining authors declare no competing interests.
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