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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 neurophysiological 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 stimulation) 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.
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 elds 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 ow 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, specicity, 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 conrmed 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 MomentumOscillatory 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:
or to G.Bák. (email:
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 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
Hz200 Hz) were used (See Box 1), while weaker intensities
(310 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 surgery810. 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
benecial 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 difcult 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 eld stimulation in brain slices, rodents, non-human
primates, and humans. We compare experimental ndings 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 ofine 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 elds
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 eld
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 rst demonstrated
this using DC stimulation applied with metal bars to the craysh
abdominal receptor and lobster cardiac ganglion neurons.
Induced electric elds as low as 1 V/m affected the timing of
action potentials while much higher elds triggered action
potentials in silent neurons23. However, the orientation of the
eld was also important: rotating the applied electric elds
orientation relative to the main neurons soma-dendritic axis
(e.g., from parallel to perpendicular to the main neuronal axis)
affected the strength and direction of the effects. These ndings
Box 1 Terminology
Transcranial electrical stimulation (TES)The technique of applying an electric eld at the scalp surface with the purpose of directly affecting brain
activity. Early efforts applied elds at high intensities for electroconvulsive therapy (ECT, 60 mA), electroanesthesia (40 mA), and electrosleep (310 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 stimulationElectrical current applied in a constant, unidirectional manner owing 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 stimulationElectrical current applied in a varying, typically sinusoidal, waveform, with current owing 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 eldThe difference in voltage between two locations (in the brain). The electric eld is a vector with both magnitude and direction, and is
measured in units of Volts/meter (V/m).
Direct effectsPotential effects of TES on the excitability of neurons in the brain, which include both immediate and cumulative effects.
Indirect effectsPotential side effects of TES on non-neuronal elements, including placebo, peripheral nerves, retina, cochlea, glia, immune system, and
blood ow as well as cumulative, long-term metabolic effects on neurons.
Immediate effectsPotential effects of TES which occur simultaneously or acutely during stimulation. These effects are measured by changes in
neuronal ring patterns or the local eld potential (dened below). We use the terms acute, concurrent, or online as synonyms of the term immediate.
Ofine effectsPotential effects of TES which may outlast the period of stimulation.
Local eld potentiala 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 eld (a vector, V/m). The local eld is monitored by small intracranial
electrodes as opposed to those obtained by scalp EEG recordings.
OscillationsRhythmic uctuations in the local eld potential, which can range in frequency from ultraslow (0.05 Hz) to ultrafast (500 Hz), with
specic oscillation frequencies characterizing particular brain states.
ShuntingDuring TES application, the ow 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.
2NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |
generalize to the mammalian brain. An anode on the pial surface
(cathode at depth) depolarizes neurons and increases neuronal
ring frequency, while the reversed current ow (surface cathode,
depth anode) hyperpolarizes neurons and reduces their ring
Yet, the above picture is an oversimplication. 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 elds. 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-specic manner without
generating action potentials. Second, although elds perpendi-
cular to the soma-dendritic axis may have little inuence on the
apical dendrites of pyramidal neurons2628, they can activate
basal dendrites and dense local axonal arbors of basket and
chandelier interneurons. In addition, neurons with star-shaped
dendritesmany inhibitory interneurons, layer 4 stellate cells,
thalamocortical cells, and basal ganglia neuronsmight be
affected equally by all eld orientations. These complexities
(Fig. 1) are relevant because the orientation of TES-induced
intracranial elds relative to these different neuronal populations
and compartments varies considerably29. In summary, the
inuence of TES current depends not only on the amplitude but
upon how the induced electric eld 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 elds on neural
We dene immediate effects of TES as those that occur simul-
taneously or acutely from the modulation of the membrane
potential by the external eld. For a sufciently strong electric
eld, neuronal ring rates should result from changes in a neu-
rons input/output function19, facilitated neurotransmitter release
from presynaptic terminals32, or ectopically induced spikes33.
With AC stimulation, current ows from anode to cathode in
one half-cycle and reverses direction during the other half-cycle.
These half-cycles elds 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
brains native oscillatory patterns3638 and thereby target cogni-
tive and therapeutic applications. While the magnitude of TES-
induced physiological effects varies with the strength of the
induced eld, the effectivenessof the stimulation and its con-
sequences on neuronal and network activity depend on many
other experimental variables and the brains state. We distinguish
ve 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 ve mechanisms of eld-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 thresholdof the induced
electric eld39,40. When a neuron nears the threshold of spike
generation, a very small amount of applied eld can bias spike
timing or spike probability. This is known as stochastic reso-
nance41. Weak stochastic effects are hard to quantify in normal
networks with uctuating activity levels because the affected
neurons may be scattered in distant networks. Upon repeated
application of the eld, different neurons may be excited or
hyperpolarized, and thus it may be difcult to generate repro-
ducible effects based on this mechanism alone. (2) Rhythm
resonance. Under well-controlled, closed-loop conditions, a very
weak eld can be precisely timed to the depolarizing phase of an
oscillating neurons membrane potential (see below). Even
without a closed-loop system, rhythm resonance may occur when
an AC eld is applied to a regular endogenous rhythm at
the same frequency, because the external eld can affect the
native oscillation at a similar phase during each cycle4250. 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 elds 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 eld. The electric eld can affect the soma,
dendrites, axon initial segment and the axon tree differently. The
relationship between the electric eld 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)
Temporal bias
of spikes
Fig. 2 Five postulated mechanisms to affect online spiking of neurons and
networks patterns in response to different estimated magnitudes of TES.
While the gure 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
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 3
forced eld may occur without affecting the overall ring 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
eld to overcome the endogenous control of network neurons
(Fig. 2).
Spike and local eld 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 difcult with current methods due to the difculty of separating
the artifact from the entrained physiological pattern, as discussed
in more detail below.
Acute effects of electric elds in vitro
In brain slices, the effects of induced elds can be studied with
exquisite control, since all synaptic signaling can be blocked
pharmacologically55. Most studies apply constant elds 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 elds of 1 V/m can polarize the soma of pyramidal neu-
rons by ~0.2 mV.
Yet, extracellular electric elds abound in the brain22. Local
electric elds (> 0.5 V/m) applied in vitro with a micropipette
near layer 5 pyramidal cell dendrites and soma induce small
intracellular voltage uctuations (e.g., 0.2 mV) and may phase-
lock the spikes to the low-frequency (130 Hz) oscillating external
eld without affecting overall spiking rate56. Thus, under idea-
lized conditions, very weak but localized extracellular elds can
bias neuronal activity (Fig. 3). Spike timing is highly sensitive to
small, persistent oscillations that act throughout a volume. This
voltage uctuation is comparable to intrinsic, non-synaptic,
internally generated noise in layer 5 pyramidal neurons (0.20.4
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 elds across the CA1 pyramidal
layer, while hippocampal sharp waves induce 515 V/m39.
Endogenous eld strengths associated with slow oscillations in
the neocortex are 12 V/m45. Whether weak applied electric elds
can overcome such strong endogenous elds remains an impor-
tant experimental question.
AC elds of sufcient strength applied at low frequencies
(< 10 Hz) can modulate ring rate periodically by the same
mechanisms as DC elds59. 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 eld, because the membrane cannot follow the
rapid uctuations in the electric eld60. The oscillatory elds may
Extracellular stimulation
Spike field
coherence (%)
Spike field
coherence (%)
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
50 mV
50 µm
1 mV
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)
uctuations (magenta). An extracellular stimulation electrode (S1) was placed 5080 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 eld 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 signicance of the spike-eld difference between control and extracellular stimulation.
dSpike-spike coordination (spike-eld coherence for two simultaneously occurring spikes; essentially spike synchrony among neurons) during 1
Hz extracellular stimulation (black) and control (cyan) condition. Note that stronger elds are needed for coherent entrainment of neuronal spikes across
the four neurons (d) than for inducing spike-eld coherence for each neuron separately (c). Figure reproduced with permission56
4NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |
not affect the overall ring 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 cycles43,45,5962. Under such ideal condi-
tions, AC elds as low as 0.20.5 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 eld at 1
Hz and > 5 V/m at 30 Hz56,39, Fig. 3), due to the frequency l-
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
and spiking
activity. These experiments have also claried the geometric
relationship between the vectorial eld 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 uid
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 elds in rodents
In rodent experiments, electric elds are often applied with
electrodes placed on the skull, the dura mater, or directly on the
brain surface. Often, a saline-lled 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 eld
and spiking rate24,34,35,40, transmembrane potential and, at
higher stimulation intensities, on the LFP25,35,39. Applied electric
elds can alter neuronal excitability64 and evoked responses64,65
in a polarity-specic 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 elds 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 eld sufcient 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.Highereld 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 (1.52 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 eld strengths across studies and even across individuals.
Opitz et al.69 measured eld strengths in two anesthetized cebus
monkeys with small round scalp electrodes (3.14 cm2) over the
left occipital and frontal cortices stimulated with 12 mA. The
median induced electric elds 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
reected the larger head musculature that shunted much of the
applied current.
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 nite-element model75, 2 mA anodal stimulation applied to the
right prefrontal region induced eld strengths between 0.40.7 V/m.
Although tDCS did not affect the spontaneous or task-related ring
rates of isolated single units or multi-unit clusters, LFP power in the
115 Hz band increased ten-fold during DC stimulation, and
coherence within the prefrontal array increased at all frequencies
from 2100 Hz. However, the changes in spike frequency or spike-
eld coherence in areas with stronger elds (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 eld 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 50120 mA (~3555 V/m electric elds
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 elds 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 eld strengths to those in rodent brains. These differences
highlight the critical need to match intracerebral elds 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 efcacy of neurons depends
on how polarized neurons align with the direction of the electrical
eld, 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 eld distribution can be obtained and
multiple stimulation parameters can be tested in the same cadaver
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 5
(Fig. 4b, c). Such empirical measurements help determine the
basic biophysical tissue properties and help model effects of TES
into the brain parenchyma7981 (Fig. 4df). Intracerebral electric
elds 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 uid behave as ohmic conductors. In other words,
induced elds 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 eld 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 eld reaching
the brain when using conventional TES, applied transcutaneously
at peak intensity currents of 12 mA, because eld strength
constrains the expected nature of the physiological changes
(Fig. 2). The main barrier for TES to reach the brain is the skulls
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 elds at the cortical surface under the electrodes
were < 0.5 V/m for 1 mA peak intensity currents68,69,84.
Can such weak electric elds 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 specic
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 rhythms37,38,62,8689, such as gamma60,89,90, beta91,
alpha46, theta92, and slow frequency oscillations40,45,93.
Whether the weak electric elds resulting from TES at con-
ventional intensities (12 mA peak) can entrain neuronal oscil-
lations is technically difcult 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 difcult. 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
ab c
de f
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 elds 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 xed 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 eld (blue open circle). fAttenuation of charge ow (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
6NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |
saturation of recording ampliers, 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 (270 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 (1014 Hz) and
high gamma (5270 Hz) frequency stimulation. These occasional
behavioral effects of TES may reect 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 thetatACS 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 nd 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 ndings 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
inuence 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 eld: 0.43 V/m). During NREM
sleep, slow oscillations (~1 Hz) strongly entrain spindle activity in
the 1015 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 4.56 mA. The effect was specic 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 12mA
(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 specic
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 elds 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 specic brain areas100.
Two other approaches attempt to increase the brain to scalp
stimulation ratio. The rst 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 ring. 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 elds 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 renements of these
techniques are necessary.
The application of high-frequency pulses avoids the saturation
of recording ampliers and allows for the simultaneous mea-
surement of the direct physiological effects of tES in humans35.
Amplier modications, 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
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 7
consistent clinical benets. 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 elds 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,broblasts112 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 eld 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 elds in the human brain. This
is sufcient to generate 0.10.2 mV changes in the membrane
potential of cells within the stimulated area. As these changes are
signicantly 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 elds 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 elds but other applications may require
higher eld strengths. For example, prompt control of spiking
activity (e.g., to terminate a seizure) may require eld strengths >
5 V/m72.
Modeling for spatial targeting and to increase replication
across subjects and studies. Individual variations in neuroa-
natomy (scalp, muscle, and skull thickness) can signicantly alter
the strength and distribution of the induced electrical eld in
humans68,98,115117 and experimental animals75,77, and thus may
account for inconsistent behavioral results in human TES
experiments. This is exemplied 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 uid volume in posterior regions lead to dif-
ferent eld distributions (Fig. 5a, b).
Computational models of current ow have been continuously
rened over the last decade78,100,116,117,120122. 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 uid. 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 eld
intensities when adjacent to CSF-lled 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 dene eld distribution for this
purpose. When an individuals 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 eld 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 F8or over left dorsolateral PFC) and current
intensities (e.g., 2mA). These specications describe how the
experiment was designed and performed, but they only indirectly
dene the induced electrical eld strength and distribution.
Reporting induced eld estimates from head models would help
relate behavioral and clinical reports to underlying neural
mechanisms. When the individual MRI head is available,
subject-specic estimates of the induced eld 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 eld strength
are included in some recent human and large animal studies but
are missing from most rodent studies. Reporting eld magnitudes
is important for animal studies where variable head size and
anatomy can strongly affect eld 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 (1100 µV). When brain
activity and stimulation are recorded simultaneously, special care
should be taken to (1) avoid saturation of the recording amplier
during experiments, and (2) subtract stimulation artifacts from
the measured signal. Saturation can be avoided by using high-
quality ampliers with sufcient input dynamic range (e.g. an
amplier with dynamic range of ± 100 mV can capture the full
stimulation artifact).
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
difcult 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.
8NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |
DC stimulation artifacts pose less of an issue during signal
processing as hardware lters will typically remove any DC offset
(avoiding amplier saturation). However, EEG recordings in
humans are affected by artifacts in the 115 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., 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
across studies.
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 benets. While the authors of
this review share a consensus on experimental recommendations,
we acknowledge a controversy within the TES eld and amongst
ourselves as to whether the attenuated electrical elds 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 eld with
competing or cooperating network oscillations when comparing
in vitro to in vivo studies; (2) external elds 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 elds 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,
Injected current
Electric field
ab c
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 conguration. cModel of Subject 2 with
another electrode conguration attempting to achieve more focal stimulation. Note that stimulation is more focal but achieves weaker elds 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-ow models were
generated from previously published data118 and methods119. We used ROAST, a toolbox for realistic current-ow models of the human119
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 9
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
desired targets.
Novel stimulation methods that safely induce higher brain to
scalp electric eld 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 efcacy of TES via resonant mod-
ulation of native brain rhythms. In selected cases, supported by
experimental ndings and human head-based modeling, sub-
cutaneous arrays or electrode plates inserted below the skin may
provide superior efcacy, 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
1. von Helmholtz, H. Treatise on Physiological Optics.Vol. III, (Dover
Publications, New York, 1867).
2. Rohracher, H. Ueber subjektive Lichterscheinungen bei Reizung mit
Wechselströmen. Zeitschrift für Sinnesphysiologie 66, 164181 (1935).
3. Barlow, H. B., Kohn, H. I. & Walsh, E. G. Visual sensations aroused by
magnetic elds. Am. J. Physiol. 148, 372375 (1947).
4. Kar, K. & Krekelberg, B. Transcranial electrical stimulation over visual cortex
evokes phosphenes with a retinal origin. J. Neurophysiol. 108, 21732178 (2012).
5. Rudorfer, M., Henry, M. & Sackeim, H. in Psychiatry (eds. Tasman, A., Kay, J.
& Lieberman, J.) 1539 (WB Saunders, Philadelphia, 1997).
6. Brown, C. C. Electroanesthesia and electrosleep. Am. Psychol. 30, 402410
7. Kuzin, M. I., Sachkov, V. I. & Zhukovskiy, V. D. Electronarcosis produced by
interference currents in clinical practice. Present-Day Technique in Surgery:
Data of the Sixth Scientic Session of the Scientic Research Institute of
Experimental Surgical Apparatus and Instruments. JPRS 31, 4,5 (1965).
8. Calancie, B. intraoperative neuromonitoring and alarm criteria for judging
MEP responses to transcranial electric stimulation: the threshold-level
method. Clin. Neurophysiol. 34,1221 (2017).
9. Merton, P. A. & Morton, H. B. Stimulation of the cerebral cortex in the intact
human subject. Nature 285, 227 (1980).
10. Rothwell, J., Thompson, P., Day, B., Boyd, S. & Marsden, C. Stimulation of the
human motor cortex through the scalp. Exp. Physiol. 76, 159200 (1991).
11. Lang, J., Sances, A. & Larson, S. J. Determination of specic cerebral
impedance and cerebral current density during the application of diffuse
electrical currents. Med. Biol. Eng. 7, 517526 (1969).
12. Toleikis, J. R., Sances, A. Jr. & Larson, S. J. Effects of diffuse transcerebral
electrical currents on cortical unit potential activity. Anesth. Analg. 53,4855
13. Nitsche, M. A. & Paulus, W. Excitability changes induced in the human motor
cortex by weak transcranial direct current stimulation. J. Physiol. 527(Pt 3),
633639 (2000).
14. Nitsche, M. A. & Paulus, W. Sustained excitability elevations induced by
transcranial DC motor cortex stimulation in humans. Neurology 57,
18991901 (2001).
15. Fry, F. J., Ades, H. W. & Fry, W. J. Production of reversible changes in the
central nervous system by ultrasound. Science 127,8384 (1958).
16. Kasschau, M. et al. Transcranial Direct Current Stimulation Is Feasible for
Remotely Supervised Home Delivery in Multiple Sclerosis. Neuromodulation
19, 824831 (2016).
17. Liu, A., Fregni, F., Hummel, F. & Pascual-Leone, A. in Transcranial Brain
Stimulation (eds. Miniussi, C., Paulus, W. & Rossini, P. M.). 359-414 (CRC
Press, Boca Raton, 2012).
18. Horvath, J. C. Are current blinding methods for transcranial direct current
stimulation (tDCS) effective in healthy populations? Clin. Neurophysiol. 126,
20452046 (2015).
19. Lafon, B. et al. Low frequency transcranial electrical stimulation does not
entrain sleep rhythms measured by human intracranial recordings. Nat.
Commun. 8, 1199 (2017).
20. Polania, R., Nitsche, M. A. & Ruff, C. C. Studying and modifying brain
function with non-invasive brain stimulation. Nat. Neurosci. 21, 174187
21. Stagg, C. J., Antal, A. & Nitsche, M. A. Physiology of transcranial direct
current stimulation. J. Ect. 34, 144152 (2018).
22. Buzsaki, G., Anastassiou, C. A. & Koch, C. The origin of extracellular elds
and currents-EEG, ECoG, LFP and spikes. Nat. Rev. Neurosci. 13, 407420
23. Terzuolo, C. A. & Bullock, T. H. Measurement of Imposed Voltage Gradient
Adequate to Modulate Neuronal Firing. Proc. Natl Acad. Sci. USA 42, 687694
24. Creutzfeldt, O. D., Fromm, G. H. & Kapp, H. Inuence of transcortical d-c
currents on cortical neuronal activity. Exp. Neurol. 5, 436452 (1962).
25. Purpura, D. P. & McMurtry, J. G. Intracellular activities and evoked potential
changes during polarization of motor cortex. J. Neurophysiol. 28, 166185
26. Ranck, J. B. Jr. Which elements are excited in electrical stimulation of
mammalian central nervous system: a review. Brain Res. 98, 417440
27. Bikson, M., Radman, T. & Datta, A. Rational modulation of neuronal
processing with applied electric elds. Conf. Proc. Ieee. Eng. Med. Biol. Soc. 1,
16161619 (2006).
28. Radman, T., Ramos, R. L., Brumberg, J. C. & Bikson, M. Role of cortical cell
type and morphology in subthreshold and suprathreshold uniform electric
eld stimulation in vitro. Brain Stimul. 2, 215228 (2009).
29. Rahman, A. et al. Cellular effects of acute direct current stimulation: somatic
and synaptic terminal effects. J. Physiol. 591, 25632578 (2013).
30. Chan, C. Y. & Nicholson, C. Modulation by applied electric elds of Purkinje
and stellate cell activity in the isolated turtle cerebellum. J. Physiol. 371,
89114 (1986).
31. Katz, B. & Schmitt, O. H. Electric interaction between two adjacent nerve
bres. J. Physiol. 97, 471488 (1940).
32. Stagg, C. J. et al. Polarity-sensitive modulation of cortical neurotransmitters by
transcranial stimulation. J. Neurosci. 29, 52025206 (2009).
33. Gutnick, M. J. & Prince, D. A. Proceedings: spontaneous antidromic spikes
from axons in cortical penicillin foci. Epilepsia 13, 354355 (1972).
34. Grossman, N. et al. Noninvasive deep brain stimulation via temporally
interfering electric elds. Cell 169, 10291041 (2017).
35. Voroslakos, M. et al. Direct effects of transcranial electric stimulation on brain
circuits in rats and humans. Nat. Commun. 9, 483 (2018).
36. Antal, A. et al. Comparatively weak after-effects of transcranial alternating
current stimulation (tACS) on cortical excitability in humans. Brain Stimul. 1,
97105 (2008).
37. Bergmann, T. O. et al. Acute changes in motor cortical excitability during slow
oscillatory and constant anodal transcranial direct current stimulation. J.
Neurophysiol. 102, 23032311 (2009).
38. Feurra, M. et al. State-dependent effects of transcranial oscillatory currents on
the motor system: what you think matters. J. Neurosci. 33, 1748317489
39. Anastassiou, C. A., Montgomery, S. M., Barahona, M., Buzsaki, G. & Koch, C.
The effect of spatially inhomogeneous extracellular electric elds on neurons.
J. Neurosci. 30, 19251936 (2010).
40. Ozen, S. et al. Transcranial electric stimulation entrains cortical neuronal
populations in rats. J. Neurosci.30, 1147611485 (2010).
41. Geisler, C. D. & Goldberg, J. M. A stochastic model of the repetitive activity of
neurons. Biophys. J. 6,5369 (1966).
42. Schutter, D. J. & Hortensius, R. Retinal origin of phosphenes to transcranial
alternating current stimulation. Clin. Neurophysiol. 121, 10801084
43. Francis, J. T., Gluckman, B. J. & Schiff, S. J. Sensitivity of neurons to weak
electric elds. J. Neurosci. 23, 72557261 (2003).
44. Rosanova, M. et al. Natural frequencies of human corticothalamic circuits.
J. Neurosci. 29, 76797685 (2009).
45. Frohlich, F. & McCormick, D. A. Endogenous electric elds may guide
neocortical network activity. Neuron 67, 129143 (2010).
46. Zaehle, T., Rach, S. & Herrmann, C. S. Transcranial alternating current
stimulation enhances individual alpha activity in human EEG. PLoS ONE 5,
e13766 (2010).
10 NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |
47. Romei, V. et al. Causal evidence that intrinsic beta-frequency is relevant for
enhanced signal propagation in the motor system as shown through rhythmic
TMS. Neuroimage 126, 120130 (2016).
48. Kanai, R., Chaieb, L., Antal, A., Walsh, V. & Paulus, W. Frequency-dependent
electrical stimulation of the visual cortex. Curr. Biol. 18, 18391843 (2008).
49. Feurra, M. et al. Frequency-dependent tuning of the human motor system
induced by transcranial oscillatory potentials. J. Neurosci. 31, 1216512170
50. Riecke, L., Formisano, E., Herrmann, C. S. & Sack, A. T. 4-Hz Transcranial
alternating current stimulation phase modulates hearing. Brain Stimul. 8,
777783 (2015).
51. Guerra, A. et al. Phase dependency of the human primary motor cortex and
cholinergic inhibition cancelation during beta tACS. Cereb. Cortex 26,
39773990 (2016).
52. Mitzdorf, U. Current source-density method and application in cat cerebral
cortex: investigation of evoked potentials and EEG phenomena. Physiol. Rev.
65,37100 (1985).
53. Nicholson, C. Theoretical Analysis of Field Potentials in Anisotropic
Ensembles of Neuronal Elements. IEEE Trans. Biomed. Eng. BME 20, 278288
54. Berenyi, A., Belluscio, M., Mao, D. & Buzsaki, G. Closed-loop control of
epilepsy by transcranial electrical stimulation. Science 337, 735737 (2012).
55. Jefferys, J. G. Inuence of electric elds on the excitability of granule cells in
guinea-pig hippocampal slices. J. Physiol. 319, 143152 (1981).
56. Anastassiou, C. A., Perin, R., Markram, H. & Koch, C. Ephaptic coupling of
cortical neurons. Nat. Neurosci. 14, 217223 (2011).
57. Jacobson, G. A. et al. Subthreshold voltage noise of rat neocortical pyramidal
neurones. J. Physiol. 564, 145160 (2005).
58. Nunez, A., Amzica, F. & Steriade, M. Electrophysiology of cat association
cortical cells in vivo: intrinsic properties and synaptic responses. J.
Neurophysiol. 70, 418430 (1993).
59. Reato, D., Rahman, A., Bikson, M. & Parra, L. C. Low-intensity electrical
stimulation affects network dynamics by modulating population rate and
spike timing. J. Neurosci. 30, 1506715079 (2010).
60. Deans, J. K., Powell, A. D. & Jefferys, J. G. Sensitivity of coherent oscillations
in rat hippocampus to AC electric elds. J. Physiol. 583, 555565 (2007).
61. Herrmann, C. S., Murray, M. M., Ionta, S., Hutt, A. & Lefebvre, J. Shaping
intrinsic neural oscillations with periodic stimulation. J. Neurosci. 36,
53285337 (2016).
62. Thut, G., Miniussi, C. & Gross, J. The functional importance of rhythmic
activity in the brain. Curr. Biol. 22, R658R663 (2012).
63. Vaidya, S. P. & Johnston, D. Temporal synchrony and gamma-to-theta power
conversion in the dendrites of CA1 pyramidal neurons. Nat. Neurosci. 16,
18121820 (2013).
64. Bindman, L. J., Lippold, O. C. & Redfearn, J. W. Long-lasting changes in the
level of the electrical activity of the cerebral cortex produced bypolarizing
currents. Nature 196, 584585 (1962).
65. Reato, D., Rahman, A., Bikson, M. & Parra, L. Effects of weak transcranial
alternating current stimulation on brain activitya review of known
mechanisms from animal studies. Front. Hum. Neurosci.7, 687 (2013).
66. Esmaeilpour, Z. et al. Notes on human trials of transcranial direct current
stimulation between 1960 and 1998. Front. Hum. Neurosci. 11, 71 (2017).
67. Jackson, M. P. et al. Animal models of transcranial direct current stimulation:
Methods and mechanisms. Clin. Neurophysiol. 127, 34253454 (2016).
68. Huang, Y. et al. Measurements and models of electric elds in the in vivo
human brain during transcranial electric stimulation. eLife 6, e18834 (2017).
69. Opitz, A. et al. Spatiotemporal structure of intracranial electric elds induced
by transcranial electric stimulation in humans and nonhuman primates. Sci.
Rep. 6, 31236 (2016).
70. Greenberg, A. & Dickson, C. T. Spontaneous and electrically modulated
spatiotemporal dynamics of the neocortical slow oscillation and associated
local fast activity. Neuroimage 83, 782794 (2013).
71. Greenberg, A., Whitten, T. A. & Dickson, C. T. Stimulating forebrain
communications: slow sinusoidal electric elds over frontal cortices
dynamically modulate hippocampal activity and cortico-hippocampal
interplay during slow-wave states. Neuroimage 133, 189206 (2016).
72. Kozak, G. & Berenyi, A. Sustained efcacy of closed loop electrical stimulation
for long-term treatment of absence epilepsy in rats. Sci. Rep. 7, 6300 (2017).
73. Greenberg, A., Abadchi, J. K., Dickson, C. T. & Mohajerani, M. H. New waves:
rhythmic electrical eld stimulation systematically alters spontaneous slow
dynamics across mouse neocortex. Neuroimage 174, 328339 (2018).
74. Krause, M. R. et al. Transcranial direct current stimulation facilitates
associative learning and alters functional connectivity in the primate brain.
Curr. Biol. 27, 30863096 (2017).
75. Datta, A. et al. On comparing in vivo intracranial recordings in non-human
primates to predictions of optimized transcranial electrical stimulation. Conf.
Proc. Ieee. Eng. Med. Biol. Soc. 2016, 17741777 (2016).
76. Kar, K., Duijnhouwer, J. & Krekelberg, B. Transcranial alternating current
stimulation attenuates neuronal adaptation. J. Neurosci. 37, 23252335 (2017).
77. Lee, W. H., Lisanby, S. H., Laine, A. F. & Peterchev, A. V. Electric eld model
of transcranial electric stimulation in nonhuman primates: correspondence to
individual motor threshold. Ieee. Trans. Biomed. Eng. 62, 20952105 (2015).
78. Datta, A. et al. Gyri-precise head model of transcranial direct current
stimulation: improved spatial focality using a ring electrode versus
conventional rectangular pad. Brain Stimul. 2, 201-207207 (2009).
79. Goncalves, S., de Munck, J. C., Verbunt, J. P., Heethaar, R. M. & da Silva, F. H.
In vivo measurement of the brain and skull resistivities using an EIT-based
method and the combined analysis of SEF/SEP data. Ieee. Trans. Biomed. Eng.
50, 11241128 (2003).
80. Haueisen, J., Ramon, C., Eiselt, M., Brauer, H. & Nowak, H. Inuence of tissue
resistivities on neuromagnetic elds and electric potentials studied with a
nite element model of the head. Ieee. Trans. Biomed. Eng. 44, 727735
81. Rush, S. & Driscoll, D. A. Current distribution in the brain from surface
electrodes. Anesth. Analg. 47, 717723 (1968).
82. Logothetis, N. K., Kayser, C. & Oeltermann, A. In vivo measurement of
cortical impedance spectrum in monkeys: implications for signal propagation.
Neuron 55, 809823 (2007).
83. Opitz, A., Falchier, A., Linn, G. S., Milham, M. P. & Schroeder, C. E.
Limitations of ex vivo measurements for in vivo neuroscience. Proc. Natl
Acad. Sci. USA 114, 52435246 (2017).
84. Chhatbar, P. Y. et al. Evidence of transcranial direct current stimulation-
generated electric elds at subthalamic level in human brain in vivo. Brain
Stimul. 11, 727733 (2018).
85. Buzsáki, G. Rhythms of the Brain (Oxford University Press, Oxford, 2006).
86. Antal, A. & Paulus, W. Transcranial alternating current stimulation (tACS).
Front. Hum. Neurosci. 7, 317 (2013).
87. Helfrich, R. F. et al. Entrainment of brain oscillations by transcranial
alternating current stimulation. Curr. Biol. 24, 333339 (2014).
88. Neuling, T., Rach, S. & Herrmann, C. S. Orchestrating neuronal networks:
sustained after-effects of transcranial alternating current stimulation depend
upon brain states. Front. Hum. Neurosci. 7, 161 (2013).
89. Reato, D. et al. Transcranial electrical stimulation accelerates human sleep
homeostasis. PLoS. Comput. Biol. 9, e1002898 (2013).
90. Voss, U. et al. Induction of self awareness in dreams through frontal low
current stimulation of gamma activity. Nat. Neurosci. 17, 810812 (2014).
91. Pogosyan, A., Gaynor, L. D., Eusebio, A. & Brown, P. Boosting cortical activity
at Beta-band frequencies slows movement in humans. Curr. Biol. 19,
16371641 (2009).
92. Marshall, L., Kirov, R., Brade, J., Molle, M. & Born, J. Transcranial electrical
currents to probe EEG brain rhythms and memory consolidation during sleep
in humans. PLoS ONE 6, e16905 (2011).
93. Marshall, L., Helgadottir, H., Molle, M. & Born, J. Boosting slow oscillations
during sleep potentiates memory. Nature 444, 610613 (2006).
94. Noury, N., Hipp, J. F. & Siegel, M. Physiological processes non-linearly affect
electrophysiological recordings during transcranial electric stimulation.
Neuroimage 140,99109 (2016).
95. Weinrich, C. A. et al. Modulation of long-range connectivity patterns via
frequency-specic stimulation of human cortex. Curr. Biol. 27, 30613068
e3063 (2017).
96. Dehaene, S., Changeux, J. P., Naccache, L., Sackur, J. & Sergent, C. Conscious,
preconscious, and subliminal processing: a testable taxonomy. Trends Cogn.
Sci. 10, 204211 (2006).
97. Williams, P. E., Mechler, F., Gordon, J., Shapley, R. & Hawken, M. J.
Entrainment to video displays in primary visual cortex of macaque and
humans. J. Neurosci. 24, 82788288 (2004).
98. Datta, A., Bikson, M. & Fregni, F. Transcranial direct current stimulation in
patients with skull defects and skull plates: high-resolution computational
FEM study of factors altering cortical current ow. Neuroimage 52, 12681278
99. Poreisz, C., Boros, K., Antal, A. & Paulus, W. Safety aspects of transcranial
direct current stimulation concerning healthy subjects and patients. Brain Res.
Bull. 72, 208214 (2007).
100. Dmochowski, J. P., Datta, A., Bikson, M., Su, Y. & Parra, L. C. Optimized
multi-electrode stimulation increases focality and intensity at target. J. Neural
Eng. 8, 046011 (2011).
101. Guleyupoglu, B., Schestatsky, P., Edwards, D., Fregni, F. & Bikson, M.
Classication of methods in transcranial electrical stimulation (tES) and
evolving strategy from historical approaches to contemporary innovations. J.
Neurosci. Methods 219, 297311 (2013).
102. Goats, G. C. Interferential current therapy. Br. J. Sports Med. 24,8792 (1990).
103. Fregni, F. et al. A sham-controlled, phase II trial of transcranial direct current
stimulation for the treatment of central pain in traumatic spinal cord injury.
Pain 122, 197209 (2006).
NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 | 11
104. Lefebvre, S. et al. Dual-tDCS enhances online motor skill learning and long-
term retention in chronic stroke patients. Front. Hum. Neurosci. 6, 343 (2012).
105. Dehaene, S. & Changeux, J. P. Experimental and theoretical approaches to
conscious processing. Neuron 70, 200227 (2011).
106. Dagan, M. et al. Multitarget transcranial direct current stimulation for freezing
of gait in Parkinsons disease. Mov. Disord. 33, 642646 (2018).
107. Fritsch, B. et al. Direct current stimulation promotes BDNF-dependent
synaptic plasticity: potential implications for motor learning. Neuron 66,
198204 (2010).
108. Johnstone, A., Levenstein, J. M., Hinson, E. L. & Stagg, C. J. Neurochemical
changes underpinning the development of adjunct therapies in recovery after
stroke: A role for GABA? J. Cereb. Blood Flow Metab. 38, 15641583 (2017).
109. Monai, H. et al. Calcium imaging reveals glial involvement in transcranial
direct current stimulation-induced plasticity in mouse brain. Nat. Commun. 7,
11100 (2016).
110. Ruohonen, J. & Karhu, J. tDCS possibly stimulates glial cells. Clin.
Neurophysiol. 123, 20062009 (2012).
111. Walleczek, J. Electromagnetic eld effects on cells of the immune system: the
role of calcium signaling. FASEB J. 6, 31773185 (1992).
112. Huang, Y. J., Samorajski, J., Kreimer, R. & Searson, P. C. The inuence of
electric eld and connement on cell motility. PLoS ONE 8, e59447 (2013).
113. Johansson, O. Disturbance of the immune system by electromagnetic elds-A
potentially underlying cause for cellular damage and tissue repair reduction
which could lead to disease and impairment. Pathophysiology 16, 157177
114. Krook-Magnuson, E., Gelinas, J. N., Soltesz, I. & Buzsaki, G. Neuroelectronics
and biooptics: closed-loop technologies in neurological disorders. JAMA
Neurol. 72, 823829 (2015).
115. Datta, A., Truong, D., Minhas, P., Parra, L. C. & Bikson, M. Inter-Individual
variation during transcranial direct current stimulation and normalization of
dose using MRI-derived computational models. Front Psychiatry.3, 91 (2012).
116. Opitz, A., Paulus, W., Will, S., Antunes, A. & Thielscher, A. Determinants of
the electric eld during transcranial direct current stimulation. Neuroimage
109, 140150 (2015).
117. Truong, D. Q., Magerowski, G., Blackburn, G. L., Bikson, M. & Alonso-
Alonso, M. Computational modeling of transcranial direct current stimulation
(tDCS) in obesity: Impact of head fat and dose guidelines. Neuroimage Clin. 2,
759766 (2013).
118. Huang, Y., Parra, L. C. & Haufe, S. The New York Head-A precise
standardized volume conductor model for EEG source localization and tES
targeting. Neuroimage 140, 150162 (2016).
119. Huang, Y., Datta A., Bikson, M. & Parra, L. Realistic volumetric-approach to
simulate transcranial electric stimulation - ROAST - a fully automated open-
source pipeline. Preprint available at (2017).
120. Miranda, P. C., Lomarev, M. & Hallett, M. Modeling the current distribution
during transcranial direct current stimulation. Clin. Neurophysiol. 117,
16231629 (2006).
121. Miranda, P. C., Mekonnen, A., Salvador, R. & Rufni, G. The electric eld in the
cortex during transcranial current stimulation. Neuroimage 70,4858 (2013).
122. Sadleir, R. J., Vannorsdall, T. D., Schretlen, D. J. & Gordon, B. Transcranial
direct current stimulation (tDCS) in a realistic head model. Neuroimage 51,
13101318 (2010).
123. Nielsen, J. D. et al. Automatic skull segmentation from MR images for realistic
volume conductor models of the head: Assessment of the state-of-the-art.
Neuroimage 174, 587598 (2018).
124. Rufni, G., Fox, M. D., Ripolles, O., Miranda, P. C. & Pascual-Leone, A.
Optimization of multifocal transcranial current stimulation for weighted
cortical pattern targeting from realistic modeling of electric elds. Neuroimage
89, 216225 (2014).
125. Noury, N. & Siegel, M. Phase properties of transcranial electrical stimulation
artifacts in electrophysiological recordings. Neuroimage 158, 406416 (2017).
126. Gebodh, N. et al. Physiologic Artifacts When Combining EEG and tDCS.
Brain Stimul. 10, e33 (2017).
127. Fertonani, A., Ferrari, C. & Miniussi, C. What do you feel if I apply
transcranial electric stimulation? Safety, sensations and secondary induced
effects. Clin. Neurophysiol. 126, 21812188 (2015).
128. Horvath, J. C., Forte, J. D. & Carter, O. Evidence that transcranial direct
current stimulation (tDCS) generates little-to-no reliable neurophysiologic
effect beyond MEP amplitude modulation in healthy human subjects: A
systematic review. Neuropsychologia 66, 213236 (2015).
129. Silberstein, S. D. et al. Safety and efcacy of peripheral nerve stimulation of the
occipital nerves for the management of chronic migraine: results from a
randomized, multicenter, double-blinded, controlled study. Cephalalgia 32,
11651179 (2012).
130. Tsubokawa, T., Katayama, Y., Yamamoto, T., Hirayama, T. & Koyama, S.
Chronic motor cortex stimulation in patients with thalamic pain. J. Neurosurg.
78, 393401 (1993).
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.
Author contributions
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.
Additional information
Supplementary Information accompanies this paper at
Competing interests: A.B. is the founder and owner of Amplipex and Neunos LLCs,
which manufacture biosignal ampliers and stimulator devices. A.B. and G.B. have led 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.
Reprints and permission information is available online at
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12 NATURE COMMUNICATIONS | (2018) 9:5092 | DOI: 10.1038/s41467-018-07233-7 |

Supplementary resource (1)

... However, they do not accurately reflect how TI-tACS acts during human neuromodulatory use, where the applied currents are much weaker. As with the past confusion about the effects of conventional tACS (Liu et al., 2018), these results highlight the importance of using realistic conditions to probe the mechanisms of human brain stimulation. Indeed, our results do seem consistent with recent experiments in humans which found that TI-tACS targeted at the hippocampus decreased BOLD activity (Violante et al., 2022). ...
Full-text available
Electrical stimulation can regulate brain activity, producing clear clinical benefits, but focal and effective neuromodulation often requires surgically implanted electrodes. Recent studies argue that temporal interference (TI) stimulation may provide similar outcomes non-invasively. During TI, scalp electrodes generate multiple electrical fields in the brain, modulating neural activity only where they overlap. Despite considerable enthusiasm for this approach, little empirical evidence demonstrates its effectiveness, especially under conditions suitable for human use. Here, using single-neuron recordings in non-human primates, we show that TI reliably alters the timing of spiking activity. However, we find that the strategies which improve the focality of TI — high frequencies, multiple electrodes, and amplitude-modulated waveforms — also limit its effectiveness. Combined, they make TI 80% weaker than other forms of non-invasive brain stimulation. Although this is too weak to cause widespread neuronal entrainment, it may be ideally suited for disrupting pathological synchronization, a hallmark of many neurological disorders.
... Evidence derived from animal models suggests that the electric field threshold for neural activation is in the range of 0.2 V m −1 -0.5 V m −1 [53][54][55][56][57]. Even electric fields as weak as 0.2 V m −1 can have an impact on neural recordings [55], and the potential mechanism of such low-intensity electrical stimulation was explained by the changes in spike probability and timing [56,58]. However, conclusions from animal models are based on brain slices or neural assemblies targeted by intracranial electrodes, and may not be directly generalizable for human studies [44]. ...
Objective. Transcranial direct current stimulation (tDCS) has been broadly used to modulate brain activity with both bipolar and high-definition montages. However, tDCS effects can be highly variable. In this work, we investigated whether the variability in the tDCS effects could be predicted by integrating individualized electric field modeling and individual pre-tDCS behavioral performance. Approach. Here, we first compared the effects of bipolar tDCS and 4 × 1 high-definition tDCS (HD-tDCS) with respect to the alleviation of visual crowding, which is the inability to identify targets in the presence of nearby flankers and considered to be an essential bottleneck of object recognition and visual awareness. We instructed subjects to perform an orientation discrimination task with both isolated and crowded targets in the periphery and measured their orientation discrimination thresholds before and after receiving 20 min of bipolar tDCS, 4 × 1 HD-tDCS, or sham stimulation over the visual cortex. Individual anatomically realistic head models were constructed to simulate tDCS-induced electric field distributions and quantify tDCS focality. Finally, a multiple linear regression model that used pre-tDCS behavioral performance and tDCS focality as factors was used to predict post-tDCS behavioral performance. Main results. We found that HD-tDCS, but not bipolar tDCS, could significantly alleviate visual crowding. Moreover, the variability in the tDCS effect could be reliably predicted by subjects' pre-tDCS behavioral performance and tDCS focality. This prediction model also performed well when generalized to other two tDCS protocols with a different electrode size or a different stimulation intensity. Significance. Our study links the variability in the tDCS-induced electric field and the pre-tDCS behavioral performance in a visual crowding task to the variability in post-tDCS performance. It provides a new approach to predicting individual tDCS effects and highlights the importance of understanding the factors that determine tDCS effectiveness while developing more robust protocols.
... The placebo effect refers specifically to changes attributed to placebo mechanisms, whereas the placebo response encompasses all health changes that occur after the administration of inactive treatment 34 . Recent research has transformed our comprehension of placebo effects from abstract to biologically based neurobehavioral phenomena 18,35 . This newfound understanding highlights their significant capability to meaningfully modulate brain regions and neurotransmitter systems. ...
... Five types of local interaction are described: stochastic resonance, rhythm resonance, temporal biasing of neuronal spikes, entrainment of network patterns, and imposed patterns (for full review see [2]). The effect of tACS can be observed during stimulation, so-called online-effect, supposedly through entrainment of EOs. ...
Full-text available
Transcranial alternating current stimulation (tACS) is a non-invasive brain stimulation method that, through its manipulation of endogenous oscillations, can affect cognition in healthy adults. Given the fact that both endogenous oscillations and cognition are impaired in various psychiatric diagnoses, tACS might represent a suitable intervention. We conducted a search of Pubmed and Web of Science databases and reviewed 27 studies where tACS is used in psychiatric diagnoses and cognition change is evaluated. TACS is a safe and well-tolerated intervention method, suitable for multiple-sessions protocols. It can be administered at home, individualized according to the patient'’s anatomical and functional characteristics, or used as a marker of disease progression. The results are varying across diagnoses and applied protocols, with some protocols showing a long-term effect. However, the overall number of studies is small with a great variety of diagnoses and tACS parameters, such as electrode montage or used frequency. Precise mechanisms of tACS interaction with pathophysiological processes are only partially described and need further research. Currently, tACS seems to be a feasible method to alleviate cognitive impairment in psychiatric patients; however, a more robust confirmation of efficacy of potential protocols is needed to introduce it into clinical practise.
... AC stimulation applied to non-active cells and neurons causes simple sinusoidal modulations of transmembrane potential that represent low-pass filtering features. Thus, high frequency tACS may be ineffective to modulate brain activity (21). However, it has been shown that network state and also the consistent AC stimulation of numerous neurons could strengthen and enhance the tACS-induced effects of polarizations. ...
Full-text available
Objective: Non-constant current stimulation (NCCS) is a neuromodulatory method in which weak alternating, pulsed or random currents are delivered to the human head via scalp or earlobe electrodes. This approach is widely used in basic and translational studies. However, the underlying mechanisms of NCCS, which lead to biological and behavioral effects in the brain, remain largely unknown. In this review, we characterize NCCS techniques currently being utilized in neuroscience investigations, including transcranial alternating current stimulation (tACS), transcranial pulsed current stimulation (tPCS), transcranial random noise stimulation (tRNS), and cranial electrotherapy stimulation (CES). Method: We unsystematically searched all relevant conference papers, journal articles, chapters, and textbooks on the biological mechanisms of NCCS techniques. Results: The fundamental idea of NCCS is that these low-level currents can interact with neuronal activity, modulate neuroplasticity and entrain cortical networks, thus, modifying cognition and behavior. We elucidate the mechanisms of action for each NCCS technique. These techniques may cause microscopic effects (such as affecting ion channels and neurotransmission systems) and macroscopic effects (such as affecting brain oscillations and functional connectivity) on the brain through different mechanisms of action (such as neural entrainment and stochastic resonance). Conclusion: The appeal of NCCS is its potential to modulate neuroplasticity noninvasively, along with the ease of use and good tolerability. Promising and interesting evidence has been reported for the capacity of NCCS to affect neural circuits and the behaviors under their control. Today, the challenge is to utilize this advancement optimally. Continuing methodological advancements with NCCS approaches will enable researchers to better understand how NCCS can be utilized for the modulation of nervous system activity and subsequent behaviors, with possible applications to non-clinical and clinical practices.
... This is the author's version which has not been fully edited and ulation signal, and the cerebral cortex is the main functional area of current signal, and regulation of the associated brain regions is realized through electrical feedback of the cerebral cortex. During the regulation process, only the amplitude and duration of signal can be changed, and the program control mode is relatively onefold [94][95][96][97]. In tACS, AC signal is used as the stimulation signal. ...
Full-text available
The sleep loss (SL) are one of the important diseases that endanger the health of aging individuals. This study evaluated SL in aging individuals, and explored the relationship between age and the organic changes that affect brain and sleep. The causes of SL in aging individuals and its harmful effects on them have been outlined. To enable individuals in choosing a suitable way to sleep better, we also reviewed advantages and disadvantages of the existing sleep modulation based on lifestyle habits and drug and physical stimulation. We found the former as more suitable for patients with mild SL, while the latter may fail to achieve the desired effects and may even lead to the onset of new diseases. Therefore, it is proposed to use non-invasive electrical and magnetic stimulations to improve the sleep quality in aging patients.In this review, mechanisms of the two stimulation methods have been summarized. Through analysis, it was found that magnetic stimulation can induce neuronal action potential, which makes patients twitch, and equipment noise causes discomfort in the elderly during treatment. Comparative analysis of stimulation methods revealed that transcranial electrical stimulation (tES) is a considerably safe, convenient, non-invasive, and easy to operate method with few adverse reactions, and it can be considered a potential therapeutic method for SL in aging patients.
Repetitive transcranial magnetic stimulation (rTMS) is a widely used therapeutic tool in neurology and psychiatry, but its cellular and molecular mechanisms are not fully understood. Standardizing stimulus parameters, specifically electric field strength and direction, is crucial in experimental and clinical settings. It enables meaningful comparisons across studies and facilitating the translation of findings into clinical practice. However, the impact of biophysical properties inherent to the stimulated neurons and networks on the outcome of rTMS protocols remains not well understood. Consequently, achieving standardization of biological effects across different brain regions and subjects poses a significant challenge. This study compared the effects of 10 Hz repetitive magnetic stimulation (rMS) in entorhino-hippocampal tissue cultures from mice and rats, providing insights into the impact of the same stimulation protocol on similar neuronal networks under standardized conditions. We observed the previously described plastic changes in excitatory and inhibitory synaptic strength of CA1 pyramidal neurons in both mouse and rat tissue cultures, but a higher stimulation intensity was required for the induction of rMS-induced synaptic plasticity in rat tissue cultures. Through systematic comparison of neuronal structural and functional properties and computational modeling, we found that morphological parameters of CA1 pyramidal neurons alone are insufficient to explain the observed differences between the groups. However, axon morphologies of individual cells played a significant role in determining activation thresholds. Notably, differences in intrinsic cellular properties were sufficient to account for the 10 % higher intensity required for the induction of synaptic plasticity in the rat tissue cultures. These findings demonstrate the critical importance of axon morphology and intrinsic cellular properties in predicting the plasticity effects of rTMS, carrying valuable implications for the development of computer models aimed at predicting and standardizing the biological effects of rTMS.
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Transcranial direct current stimulation (tDCS) has been used to noninvasively reduce epileptic activity in focal epilepsy. In this proof-of-principle N-of-1 trial in a patient with drug-resistant focal epilepsy, we propose distributed constrained maximum intensity (D-CMI) for individually targeted and optimized multi-channel (mc-)tDCS to reduce epileptic activity. Combined electro- and magnetoencephalography (EMEG) source analysis in a realistic calibrated head model defines location and orientation of the target epileptogenic source. Converging evidence for this determination is achieved by retrospective identification of a cortical malformation in magnetic resonance imaging and by successful EMEG-guided invasive EEG. We applied D-CMI in a double-blind, sham-controlled stimulation experiment. In two stimulation weeks, either D-CMI or sham stimulation with 4 mA injection current were applied twice every week-day for 20 min each, with a 20 min pause in between. EEG was recorded 1 h before and after stimulation. For D-CMI, we find a highly significant reduction in IED frequency (p < 0.0001) marked by three experts of on average 37% to 81% over the five days of stimulation (mean +- SD: 58% +- 19%), while this is not the case for sham. The proposed procedure was well-tolerated and parameterizes a group clinical trial (Study registration number: DRKS00029384).
Experimental data suggest a role of blood–brain barrier P-glycoprotein overexpression as a factor contributing to drug-resistant epilepsy. Therefore, efforts are made to develop and validate therapeutic approaches that aim to overcome transporter-mediated drug resistance. Considering the physiological function of efflux transporters, it will be advantageous to preserve the basal transport function. Thus, recent studies focused on the elucidation of key signaling factors driving P-glycoprotein upregulation in response to epileptic seizure activity. Based on these investigations, novel concepts have been developed: blocking the signaling pathway and controlling P-glycoprotein expression despite recurrent seizure activity. Further development of respective approaches is based on ongoing studies that explore the relevance of the signaling mechanisms in human capillaries and preclinical animal models. In general, the success of respective strategies will depend on the question of whether patients exist in which P-glycoprotein overexpression constitutes a predominant factor contributing to therapeutic failure.KeywordsP-glycoproteinEfflux transportersGlutamateCyclooxygenase-2Drug-resistant Epilepsy
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Research in the area of transcranial electrical stimulation (TES) often relies on computational models of current flow in the brain. Models are built on magnetic resonance images (MRI) of the human head to capture detailed individual anatomy. To simulate current flow, MRIs have to be segmented, virtual electrodes have to be placed on these anatomical models, the volume is tessellated into a mesh, and the finite element model is solved numerically to estimate the current flow. Various software tools are available for each step, as well as processing pipelines that connect these tools for automated or semi-automated processing. The goal of the present tool -- ROAST -- is to provide an end-to-end pipeline that can automatically process individual heads with realistic volumetric anatomy leveraging open-source software (SPM8, iso2mesh and getDP) and custom scripts to improve segmentation and execute electrode placement. When we compare the results on a standard head with other major commercial software tools for finite element modeling (ScanIP, Abaqus), ROAST only leads to a small difference of 9% in the estimated electric field in the brain. We obtain a larger difference of 47% when comparing results with SimNIBS, an automated pipeline that is based on surface segmentation of the head. We release ROAST as a fully automated pipeline available online as a open-source tool for TES modeling.
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Background: Recent findings suggest that transcranial direct current stimulation of the primary motor cortex may ameliorate freezing of gait. However, the effects of multitarget simultaneous stimulation of motor and cognitive networks are mostly unknown. The objective of this study was to evaluate the effects of multitarget transcranial direct current stimulation of the primary motor cortex and left dorsolateral prefrontal cortex on freezing of gait and related outcomes. Methods: Twenty patients with Parkinson's disease and freezing of gait received 20 minutes of transcranial direct current stimulation on 3 separate visits. Transcranial direct current stimulation targeted the primary motor cortex and left dorsolateral prefrontal cortex simultaneously, primary motor cortex only, or sham stimulation (order randomized and double-blinded assessments). Participants completed a freezing of gait-provoking test, the Timed Up and Go, and the Stroop test before and after each transcranial direct current stimulation session. Results: Performance on the freezing of gait-provoking test (P = 0.010), Timed Up and Go (P = 0.006), and the Stroop test (P = 0.016) improved after simultaneous stimulation of the primary motor cortex and left dorsolateral prefrontal cortex, but not after primary motor cortex only or sham stimulation. Conclusions: Transcranial direct current stimulation designed to simultaneously target motor and cognitive regions apparently induces immediate aftereffects in the brain that translate into reduced freezing of gait and improvements in executive function and mobility. © 2018 International Parkinson and Movement Disorder Society.
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Transcranial electric stimulation is a non-invasive tool that can influence brain activity; however, the parameters necessary to affect local circuits in vivo remain to be explored. Here, we report that in rodents and human cadaver brains, ~75% of scalp-applied currents are attenuated by soft tissue and skull. Using intracellular and extracellular recordings in rats, we find that at least 1 mV/mm voltage gradient is necessary to affect neuronal spiking and subthreshold currents. We designed an 'intersectional short pulse' stimulation method to inject sufficiently high current intensities into the brain, while keeping the charge density and sensation on the scalp surface relatively low. We verify the regional specificity of this novel method in rodents; in humans, we demonstrate how it affects the amplitude of simultaneously recorded EEG alpha waves. Our combined results establish that neuronal circuits are instantaneously affected by intensity currents that are higher than those used in conventional protocols.
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Transcranial electrical stimulation has widespread clinical and research applications, yet its effect on ongoing neural activity in humans is not well established. Previous reports argue that transcranial alternating current stimulation (tACS) can entrain and enhance neural rhythms related to memory, but the evidence from non-invasive recordings has remained inconclusive. Here, we measure endogenous spindle and theta activity intracranially in humans during low-frequency tACS and find no stable entrainment of spindle power during non-REM sleep, nor of theta power during resting wakefulness. As positive controls, we find robust entrainment of spindle activity to endogenous slow-wave activity in 66% of electrodes as well as entrainment to rhythmic noise-burst acoustic stimulation in 14% of electrodes. We conclude that low-frequency tACS at common stimulation intensities neither acutely modulates spindle activity during sleep nor theta activity during waking rest, likely because of the attenuated electrical fields reaching the cortical surface.
Direct current stimulation is a neuromodulatory noninvasive brain stimulation tool, which was first introduced in animal and human experiments in the 1950s, and added to the standard arsenal of methods to alter brain physiology as well as psychological, motor, and behavioral processes and clinical symptoms in neurological and psychiatric diseases about 20 years ago. In contrast to other noninvasive brain stimulation tools, such as transcranial magnetic stimulation, it does not directly induce cerebral activity, but rather alters spontaneous brain activity and excitability by subthreshold modulation of neuronal membranes. Beyond acute effects on brain functions, specific protocols are suited to induce long-lasting alterations of cortical excitability and activity, which share features with long-term potentiation and depression. These neuroplastic processes are important foundations for various cognitive functions such as learning and memory formation and are pathologically altered in numerous neurological and psychiatric diseases. This explains the increasing interest to investigate transcranial direct current stimulation (tDCS) as a therapeutic tool. However, for tDCS to be used effectively, it is crucial to be informed about physiological mechanisms of action. These have been increasingly elucidated during the last years. This review gives an overview of the current knowledge available regarding physiological mechanisms of tDCS, spanning from acute regional effects, over neuroplastic effects to its impact on cerebral networks. Although knowledge about the physiological effects of tDCS is still not complete, this might help to guide applications on a scientifically sound foundation.
Background: Transcranial direct current stimulation (tDCS) is a promising brain modulation technique for several disease conditions. With this technique, some portion of the current penetrates through the scalp to the cortex and modulates cortical excitability, but a recent human cadaver study questions the amount. This insufficient intracerebral penetration of currents may partially explain the inconsistent and mixed results in tDCS studies to date. Experimental validation of a transcranial alternating current stimulation-generated electric field (EF) in vivo has been performed on the cortical (using electrocorticography, ECoG, electrodes), subcortical (using stereo electroencephalography, SEEG, electrodes) and deeper thalamic/subthalamic levels (using DBS electrodes). However, tDCS-generated EF measurements have never been attempted. Objective: We aimed to demonstrate that tDCS generates biologically relevant EF as deep as the subthalamic level in vivo. Methods: Patients with movement disorders who have implanted deep brain stimulation (DBS) electrodes serve as a natural experimental model for thalamic/subthalamic recordings of tDCS-generated EF. We measured voltage changes from DBS electrodes and body resistance from tDCS electrodes in three subjects while applying direct current to the scalp at 2 mA and 4 mA over two tDCS montages. Results: Voltage changes at the level of deep nuclei changed proportionally with the level of applied current and varied with different tDCS montages. Conclusions: Our findings suggest that scalp-applied tDCS generates biologically relevant EF. Incorporation of these experimental results may improve finite element analysis (FEA)-based models.
Anatomically realistic volume conductor models of the human head are important for accurate forward modeling of the electric field during transcranial brain stimulation (TBS), electro- (EEG) and magnetoencephalography (MEG). In particular, the skull compartment exerts a strong influence on the field distribution due to its low conductivity, suggesting the need to represent its geometry accurately. However, automatic skull reconstruction from structural magnetic resonance (MR) images is difficult, as compact bone has a very low signal in magnetic resonance imaging (MRI). Here, we evaluate three methods for skull segmentation, namely FSL BET2, the unified segmentation routine of SPM12 with extended spatial tissue priors, and the skullfinder tool of BrainSuite. To our knowledge, this study is the first to rigorously assess the accuracy of these state-of-the-art tools by comparison with CT-based skull segmentations on a group of ten subjects. We demonstrate several key factors that improve the segmentation quality, including the use of multi-contrast MRI data, the optimization of the MR sequences and the adaptation of the parameters of the segmentation methods. We conclude that FSL and SPM12 achieve better skull segmentations than BrainSuite. The former methods obtain reasonable results for the upper part of the skull when a combination of T1- and T2-weighted images is used as input. The SPM12-based results can be improved slightly further by means of simple morphological operations to fix local defects. In contrast to FSL BET2, the SPM12-based segmentation with extended spatial tissue priors and the BrainSuite-based segmentation provide coarse reconstructions of the vertebrae, enabling the construction of volume conductor models that include the neck. We exemplarily demonstrate that the extended models enable a more accurate estimation of the electric field distribution during transcranial direct current stimulation (tDCS) for montages that involve extraencephalic electrodes. The methods provided by FSL and SPM12 are integrated into pipelines for the automatic generation of realistic head models based on tetrahedral meshes, which are distributed as part of the open-source software package SimNIBS for field calculations for transcranial brain stimulation.
In the past three decades, our understanding of brain-behavior relationships has been significantly shaped by research using non-invasive brain stimulation (NIBS) techniques. These methods allow non-invasive and safe modulation of neural processes in the healthy brain, enabling researchers to directly study how experimentally altered neural activity causally affects behavior. This unique property of NIBS methods has, on the one hand, led to groundbreaking findings on the brain basis of various aspects of behavior and has raised interest in possible clinical and practical applications of these methods. On the other hand, it has also triggered increasingly critical debates about the properties and possible limitations of these methods. In this review, we discuss these issues, clarify the challenges associated with the use of currently available NIBS techniques for basic research and practical applications, and provide recommendations for studies using NIBS techniques to establish brain-behavior relationships.
There has been growing interest in transcranial direct current stimulation (tDCS), a non-invasive technique purported to modulate neural activity via weak, externally applied electric fields. Although some promising preliminary data have been reported for applications ranging from stroke rehabilitation to cognitive enhancement, little is known about how tDCS affects the human brain, and some studies have concluded that it may have no effect at all. Here, we describe a macaque model of tDCS that allows us to simultaneously examine the effects of tDCS on brain activity and behavior. We find that applying tDCS to right prefrontal cortex improves monkeys' performance on an associative learning task. While firing rates do not change within the targeted area, tDCS does induce large low-frequency oscillations in the underlying tissue. These oscillations alter functional connectivity, both locally and between distant brain areas, and these long-range changes correlate with tDCS's effects on behavior. Together, these results are consistent with the idea that tDCS leads to widespread changes in brain activity and suggest that it may be a valuable method for cheaply and non-invasively altering functional connectivity in humans.