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DOI: 10.1038/s41467-018-02928-3

OPEN

Direct effects of transcranial electric stimulation on
brain circuits in rats and humans

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Mihály Vöröslakos 1,2, Yuichi Takeuchi 1, Kitti Brinyiczki3, Tamás Zombori3, Azahara Oliva1,2,
Antonio Fernández-Ruiz1,2, Gábor Kozák 1, Zsigmond Tamás Kincses 4, Béla Iványi3,
György Buzsáki 2,5,6 & Antal Berényi 1,2

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 areattenuated 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 andsubthreshold currents. We designed an ‘intersectional short pulse’ stimulation method to
inject sufﬁciently high current intensities into the brain, while keeping the charge density andsensation 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.

1 MTA-SZTE “Momentum” Oscillatory Neuronal Networks Research Group, Department of Physiology, University of Szeged, Szeged, 6720, Hungary. 2 The
Neuroscience Institute, New York University, New York, NY 10016, USA. 3 Department of Pathology, University of Szeged, Szeged, 6725, Hungary.
4 Department of Neurology, University of Szeged, Szeged, 6725, Hungary. 5 Department of Neurology, New York University, New York, NY 10016, USA.
6 Center for Neural Science, New York University, New York, NY 10016, USA. Correspondence and requests for materials should be addressed toG.Bák. (email: gyorgy.buzsaki@nyumc.org) or to A.Bén. (email: drberenyi@gmail.com)

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3

E

lectric ﬁelds, generated either by neurons themselves orapplied externally, can affect the transmembrane potentialof neurons and, consequently, the probability of occurrenceof action potentials1–3. Forced electric ﬁelds, induced either
locally (e.g., by deep brain stimulation) or non-invasively throughthe scalp (e.g., by transcranial electrical stimulation; TES)4–7, canprobe neural patterns and potentially to ameliorate brain disease8.
Yet, there is no accepted theory to explain how TES affects
neuronal circuits in the brain, mainly because the physiologicalmechanisms of TES are not well understood. Electrical stimulation of the scalp can affect brain activity in multiple indirect ways,
including activation of peripheral nerves7,9,10, retina11, the vestibular apparatus, astrocytes, perivascular elements12,13, andplacebo effects14. Given the important role of brain oscillations in
cognition, an often-stated explicit goal of TES is to bias brain
rhythms acutely or chronically15,16, as opposed to inducing
indirect peripheral effects.
For many therapeutic applications, it is desirable to affectneurons promptly (e.g., to terminate epileptic seizures), in aregionally constrained manner to reach maximum on-target
effects and reduce effects on unintended brain networks4,5.
Achieving spatially precise effects by scalp-applied currents
requires knowledge about the spread of electric ﬁelds in thehuman head17 and the use of novel methods of current delivered
through multiple electrodes18,19.

a

The effectiveness of currently used TES protocols on local
neuronal networks is a subject of extensive debate14–16,20,21. At
least two factors contribute to this controversy. First, the large
electric ﬁelds induced by alternating current TES (transcranialalternating current stimulation; tACS) often prevents simultaneous measurement of electric (electroencephalographic, EEG),
magnetic (magnetoencephalographic, MEG), or imaging (blood
oxygen-level dependent, BOLD) signals22. Recent experiments
attempted to alleviate the ampliﬁer saturation problem andremove the stimulus artifacts21–23. However, in those experiments
the expected brain rhythm entrainment was examined at thesame frequency of the applied TES (e.g., 10 Hz tACS induced
increased power in the alpha band), raising the possibility thatlarge tACS artifacts which are several thousand-fold larger than
the scalp signal, or a harmonic of the artifact, have contaminatedthe results. A second indirect approach takes the voltage gradientsshown experimentally to produce spike entrainment and estimates the corresponding current intensity applied at the scalp
surface. However, translation of results obtained from models24,
in vitro observations21,22,25,26, and experiments performed on
experimental animals6,27 to humans is complicated by anincomplete understanding of how skin, subcutaneous soft tissue,
skull, cerebrospinal ﬂuid, and brain folding affects current
spread28. While strong stimulation (>50 mA; 0.5 ms pulses)
delivered through intracranial screw electrodes in anesthetized

b

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Fig. 1 Intracerebral electric ﬁeld distribution and magnitude during subcutaneous or transcutaneous stimulation. a Intraoperative photograph of themeasurement of intracerebral electric ﬁelds by a 6-by-5 electrode matrix in an anesthetized rat. Red: cathodal, blue: anodal gel electrodes glued onto theskull surface. The spline interpolated map of the measured intracerebral gradients shown in b (bottom) is superimposed on the skull surface. b Map of themeasured intracerebral gradients in the horizontal plane. The boundary where the gradients are >1 mV/mm is marked by black lines (applied intensity:
100 µA). Right, boxplots with whiskers indicate group results (full data set is superimposed in gray). Proper placement of the stimulating electrodes canrestrict the extent of the effective electric ﬁeld gradients to either the left (upper plot) or the right (lower plot) hemisphere (P = 0.04 and 0.02; n = 20 in 4
rats each, Mann–Whitney U-test). c Photograph of scalp stimulation electrodes and the small hole in the skull through which intracellular recordings were
made. d, e Transcutaneous stimulation at the same stimulus intensities generated several-fold weaker electric ﬁelds compared to subcutaneous stimulation(P < 0.005, n = 20 in 4 rats)
2

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a

Skull

15
Δ Membrane potential (mV)Skin

* ** ****** *** ******* *** *
Skin
Skull

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–30
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Frequency (Hz)Fig. 2 Modulating neuronal activity by subcutaneous or transcutaneous stimulation. a Subthreshold membrane potential changes of cortical neurons bytranscutaneous and subcutaneous direct current stimuli. Vm was held below spiking by intracellularly injected hyperpolarizing current. Five representative
trials are shown for each arrangement. Right panel, group effects (n = 40 trials from 8 neurons of 3 rats for transcutaneous and n = 25 trials from 5 neuronsof 4 rats for subcutaneous experiments). Note linear changes of Vm with changing polarity and amplitude of forced ﬁelds (R = 0.86, P < 0.005 fortranscutaneous and R = 0.97, P < 0.005 for subcutaneous stimulation, n = 13 trials, each; asterisks mark signiﬁcant differences against control condition, n= 25/40 for subcutaneous/transcutaneous trials). For each stimulus intensity, the generated electric ﬁeld strengths are shown at the bottom of the plot inblue and red for transcutaneous and subcutaneous stimuli, respectively. b Same as a but for affected spiking frequency by applied ﬁelds (R = 0.80, P =0.007 for transcutaneous and R = 0.95, P < 0.005 for subcutaneous stimulation, n = 13, each; asterisks mark signiﬁcant pairwise differences against
control condition, n = 25/35 for subcutaneous/transcutaneous trials). c, d Changes of Vm power spectra in response to subcutaneous (c, n = 30 trials) andtranscutaneous (d, n = 35 trials) stimuli. Note the lack of a signiﬁcant effect with transcutaneous stimulation and prominent decrease of delta power (1–5.4
Hz) at +600 and +800 µA conditions compared to control (arrow; P < 0.005, n = 30 power value pairs at each frequency bin from 6 animals;
Mann–Whitney U-test with Bonferroni correction)

patients has shown convincing brain network-induced
effects29,30, the current intensity applied to the scalp needed toacutely affect neuronal patterns is yet to be established5,31,32.
The goal of the experiments presented in this paper was to
identify the conditions under which neuronal spikes and local
circuits could be directly affected by TES. To accomplish this goal,
we ﬁrst determined the voltage gradients necessary to affect themembrane potential and neuronal spiking in the intact rat brain,
NATURE COMMUNICATIONS | (2018)9:483

corresponding to approximately 1 mV/mm. Second, we introduced a novel fast pulse stimulation method that allowed
simultaneous recording of electrical activity and focusing induced
ﬁelds to target brain structures. We veriﬁed the validity of thismethod in rats and tested it in human subjects. Third, we measured the impact of scalp, soft tissue, and skull on current spreadand quantiﬁed the induced ﬁelds in the brain of human cadavers,
and found that only approximately 25% of the scalp-applied| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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ac

Example putative pyramidal neuron
from left hippocampus

Model of focused TES

5

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Ipsi

Time (μs)∫ (Temporal summation)

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Spikes

500 ms10
Pyramidal cell
Interneuron
No change

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ISP

Contralateral
ISP

Fig. 3 Intersectional Short Pulse (ISP) stimulation can spatially focus induced ﬁelds. a Leaky integrate-and-ﬁre neuron model cartoon to demonstrate theconcept of ISP stimulation. Stimulus current is delivered sequentially through three independent electrode pairs generating a continuously changing
intracerebral gradient pattern. Neuronal cell membranes can integrate these patterns due to their relatively slow membrane time constant (10 ms).
Consequently, neurons at the cross-section of the current ﬂow axes integrate all three stimuli, and become more strongly entrained than neurons locatedoutside the focus. b Experimental protocol to measure the efﬁcacy of ISP. White circle marks the craniotomy for the example left hippocampal neuron
shown in c. The contralateral craniotomy is not displayed for simplicity. 3D-printed gel electrode holders (anode = left; cathode = right) were attached tothe temporal bones bilaterally with ﬁve electrodes on each side. Three electrode pairs were programmed to target the ISP beams on either the left or theright hemisphere (serving as ISPipsi and ISPcontra conditions for the example in c, respectively). Each electrode pair was pulsed for 2.5 µs and the pulses
cycled through the three pairs for 500 ms followed by non-stimulated 1-s control periods. This sequence was repeated to alternatingly stimulate the rightor left hemisphere. The idealized beam crossings shown here may be modiﬁed by the inhomogeneity of brain structures and ventricles. c Response of an
example neuron. The putative pyramidal cell from the left hippocampus was strongly excited by the ipsilateral focal stimulation, as shown by peristimulus
time histograms (top panels) and raster plots (middle panels). ISP stimulation did not affect isolation of single units as demonstrated by the similar
autocorrelograms and identical spike waveforms during stimulation and control periods. d Fold-changes of normalized ﬁring rates of the signiﬁcantly
modulated cells from the left (n = 32 units) and from the right hippocampus (n = 23 units) show lateralized effect of the ISP stimulation (P = 0.001;
ISPipsilateral vs. ISPcontralateral Wilcoxon signed rank test)

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current enters the brain. Finally, we determined the current levels
of TES necessary to affect the amplitude of alpha waves in humansubjects. Our direct measurements and indirect estimation provided concordant results and established that in humans at least
4–6 mA currents should be applied by conventional tACS electrodes to reliably and instantaneously affect neuronal circuits.
Results
Subcutaneous and transcutaneous electric stimulation in rats.
Previous in vivo experiments in rodents were performed using
stimulation electrodes placed on the skull4,6 (i.e., subcutaneous
TES) and demonstrated both stimulus-locked ﬁring of neurons inboth neocortex and subcortical structures6 as well as effects on
the amplitude of intracerebrally recorded local ﬁeld potentials(LFP)4. Because skin and head musculature surrounding the skull
represent a large shunt which can diffuse the applied current, ﬁrst
we examined the current loss between scalp and the brain. Using
implanted arrays of electrodes and stimulation pads placed on theparietal bone and pre-bregma frontal bone, we measured theintracerebral voltage gradients in the horizontal plane (Fig. 1a, b).
Subcutaneous TES at 50 µA alternating current was sufﬁcient to
induce ~1 mV/mm electric ﬁelds (Fig. 1b, d). The distribution of
the electric ﬁelds could be biased by different conﬁgurations ofbipolar stimulation (Fig. 1b, left-side stimulation: 20.33 (IQR =11.49–24.52) and 3.65 (IQR = 1.58–8.39) mm2 exceeding 1 mV/
mm on the left and right hemisphere, respectively; P = 0.04; rightside stimulation: 5.16 (IQR = 1.01–21.76) and 28.37 (IQR =20.72–31.18) mm2 exceeding 1 mV/mm on the left and righthemisphere, respectively; P = 0.02; Mann–Whitney U-test and n= 20 × 2 in both cases). In contrast, applying the current to theshaved scalp (transcutaneous TES) through the same size electrodes resulted in an 80 ± 5% current loss, independent of thestimulus intensity (Fig. 1c–e; subcutaneous stimulation: 17.01(IQR = 14.96–20.85) mV/mm/mA; transcutaneous stimulation:
2.14 (IQR = 1.9–2.44) mV/mm/mA; P < 0.001; paired t-test; n =20 × 2).
In a more direct physiological comparison, we tested the effects
of externally applied direct currents on the intracellularly
recorded transmembrane potential (Vm) and spiking of neurons
in the deep layers of the visual cortex (Fig. 2a–d). Subcutaneous(skull) stimulation exerted clear and predictable effects on Vm.
Depending on the polarity of the stimulation, Vm became
depolarized or hyperpolarized in a relatively linear manner (Fig.
2a; Pearson’s linear correlation, R = 0.86, P = 0.002 for transcutaneous and R = 0.97, P < 0.001 for subcutaneous stimulation, n =13, each), and decreased or increased the number of action
potentials, respectively (Fig. 2b; Pearson’s linear correlation, R =0.80, P = 0.007 for transcutaneous and R = 0.95, P < 0.001 forsubcutaneous stimulation, n = 13, each). Subcutaneous depolarizing pulses signiﬁcantly decreased Vm (paired t-test with
Bonferroni correction; P <0.001, <0.001, <0.001 for 400, 600,
800 µA vs. 0 µA; n = 25 membrane potential difference values),
increased ﬁring rate (paired t-test with Bonferroni correction; P =0.001, <0.001, <0.001 for 200, 600, 800 µA vs. 0 µA; n = 25 ﬁring
rate difference values) and reduced Vm power in delta frequency
band (1–5 Hz; +600 to +800 µA, Mann–Whitney U-test with
Bonferroni correction; P < 0.005, n = 30 power value pairs at eachfrequency bin), indicating that subcutaneous stimulation affectedmany other neurons as well (Fig. 2c). Hyperpolarizing pulses
exerted opposite effect with similar magnitudes on Vm (paired ttest with Bonferroni correction; P = 0.003, 0.004 and 0.046 for−800, −600, −400 µA vs. 0 µA; n = 25 membrane potential
difference values) and reduced ﬁring rate (paired t-test with
Bonferroni correction; P = 0.044, 0.028 for −800 and −600 µA vs.
0 µA; n = 25 ﬁring rate difference values). Using the same current
NATURE COMMUNICATIONS | (2018)9:483

intensities, transcutaneous (scalp) stimulation produced muchsmaller and more variable effects (Vm was affected at anodal
400–800 µA but not by cathodal pulses; paired t-test with
Bonferroni correction; P = 0.044, 0.008, and 0.003 for 400, 600,
and 800 µA vs. 0 µA; n = 40 membrane potential difference values,and even the highest current intensities failed to affect delta
power Vm or higher frequencies; Fig. 2d; Mann–Whitney U-test
with Bonferroni correction; P > 0.05; n = 35 × 150 spectral amplitude values for all conditions). Spiking activity by transcutanenous stimulation was affected at only 800 µA depolarizing pulses(paired t-test with Bonferroni correction; P = 0.046 for 800 µA vs.
0 µA; n = 35 ﬁring rate difference values), corresponding to
intracranial ﬁelds of approximately 2 mV/mm (Fig. 2b, d). In
summary, electric ﬁelds applied either subcutaneously ortranscutaneously, which induce at least 1 mV/mm intracerebral
voltage gradient, can affect spiking activity, but stronger ﬁelds areneeded to affect network oscillations.

Focused TES effect by Intersectional Short Pulse stimulation.
For many experimental and clinical applications5,22, it would bedesirable to apply TES in a spatially targeted manner andsimultaneously monitor the induced electrical changes to verify
online effectiveness. Because the scalp, skull, and brain conduct
current in a homogenous manner, simultaneous application of
TES through multiple electrode pairs cannot induce a spatiallyconﬁned effect (see Supplementary Figure 6b). Our proposedsolution to achieve spatially targeted TES effects is to apply
spatio-temporally rotating Intersectional Short Pulse (ISP) stimulation. This method exploits the short integration time constant of the neuronal membrane (5–20 ms), a mechanism that cantemporally integrate multiple electrical gradients with similar
vector directions (Fig. 3a, Supplementary Figure 1)31. An added
advantage of fast pulse stimulation (2.5 or 10 µs pulse width with
5 or 50 µs pause, depending on the number of electrode pairs) isthat the transients of high frequency pulses affect simultaneously
recorded LFP or neuronal spikes (1 Hz–5 kHz; 20 kHz sampling)
substantially less than conventional tACS and they do not saturate recording ampliﬁers even at relatively high intensities.
To test our model prediction of focal effect in rats, current
pulses were delivered sequentially and in a spatially asymmetricmanner through independently programmable isolated current
generators, which were connected to a three-dimensional (3-D)
printed gel-electrode strip glued to the temporal bone surface(Fig. 3b). Unit activity was recorded bilaterally in the hippocampal CA1 region with silicon probes (7 anesthetized rats and 1
chronically implanted rat). The hemisphere target of the bipolarstimulation conﬁguration was alternated (Fig. 3b, c; Supplementary Figure 2a). The effectiveness of the ISP stimulation on
spatially targeted entrainment of single unit activity is illustrated
for an example neuron from the left hippocampus (Fig. 2c). The
artifacts of the short duration stimulation pulses did not affect therecording quality as demonstrated by the similar spike waveforms
and spike autocorrelograms of the unit during stimulation andstimulation-free periods (Fig. 3c). Of the 127 isolated single units,
55 were signiﬁcantly affected by at least one conﬁguration of thestimulation protocol (32 increased and 23 decreased, signiﬁcancethreshold: P < 0.05; Wilcoxon signed rank test; n = 300 ﬁring rate
values for each neuron and condition, all tested against baseline
condition). To quantify the focusing effect of ISP, we calculated
the fold-change of unit discharge in the left and righthippocampus, respectively. Using only three rotating dipoles,
the current-focusing effect of ISP resulted in a several-foldincrease in induced unit discharge between the targeted and nontargeted hemispheres (Fig. 3d, 1.8 ± 2.35-fold vs. 1.017 ± 0.63-| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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fold; mean ± SD; P = 0.001; Wilcoxon signed rank test; n = 55
units).
In four of the above animals (three anesthetized and one
chronic; 77 units), the spatial selectivity of the ISP method wascompared to traditional direct current (DC) pulses (for DC
stimulation, the electrodes in the same hemisphere were shortcircuited; the same current intensity was used for DC and ISP
stimulation). For each protocol, 500 ms stimulation epochsaalternated with 1000 ms stimulus-free epochs using the following
sequence: ISPleft, ISPright, DCleft, DCright. Eighteen (ISP) and 10(DC) neurons showed signiﬁcant ﬁring rate changes to at least
one stimulation conﬁguration (signiﬁcance threshold; P < 0.05;
Wilcoxon signed rank test; n = 340 ﬁring rate values for eachneuron and condition, tested against baseline). Of the 18 ISPdriven hippocampal neurons, eight (44%) responded differentially to ISPleft and ISPright conditions (signiﬁcance threshold: P <bContact site

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Stimulus frequency (Hz)Case-1Case-2

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Measuring current spread in human cadavers. Currently, thebest estimates of the currents needed to induce electric ﬁelds of
sufﬁcient magnitude intracranially are offered by in silico modeling of the human head5,33. However, there are many uncertainties in such modeling. As an alternative to modeling, we
carried out high spatial density, 3-D intracerebral measurements
in cadaver brains in situ (n = 11; Supplementary Tables 1 and 2).

Current (mA)

6 × 10 mm

c
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Skull

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Distance from skin surface (mm)Stimulationelectrode

Skin

Skull

Brain

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2 3 4 5
Voltage (V)

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1.5

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Relative shunting effect

***

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in
Sk
ul
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Normalized electric field(mV/mm/mA)

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Electric field (mV/mm)b

5 × 15 mm

Shunting effect (%)

a

Prior to each experiment, the subdural space was ﬁlled withphysiological saline (0.9% NaCl) solution to replace the cerebrospinal ﬂuid lost during the insertion of the recording electrodes. Thirty-six custom-made multisite electrodes (three toseven sites per electrode, 198 in total, Fig. 4a, Supplementary
Figure 3) were inserted into the brain through holes drilled
through the calvarium after removing the soft tissue around theskull (Fig. 4b, c) to create a 3-D montage. As the overall volumeof the removed skull was negligible compared to the total skullvolume, and the polyimide electrode shafts were tightly sealed,
the conduction/isolation properties of the skull were not affected.
A needle electrode placed into the sagittal sulcus on the forehead
served as the reference electrode. Four or seven pairs of Ag/AgCl
stimulation electrodes were ﬁxed to the skull surface bilaterally byconductive electrode gel (Fig. 4b, c). Using either DC pulses oralternating current stimulation (Supplementary Figure 4) showedElectric field (mV/mm)

0.05; paired t-test; n = 340 ﬁring rate pairs for each neuron). Of
the 10 DC-driven neurons, only one neuron (10%) showedsigniﬁcant difference to left (anode) vs. right (anode) stimulation(signiﬁcance threshold: P < 0.05; paired t-test; n = 340 ﬁring ratepairs for each neuron). In summary, the ISP stimulation affectedneural activity in spatially targeted manner, even though skull
thickness, brain geometry, tissue anisotropy, and ventricles likely
distorted current spread.

50

0
Skin

Skull

Brain

4 mm
5 mm
6 mm
7 mm

1

0.5

0

0

10
20
30
40
Distance from skull (mm)Fig. 5 Skin and subcutaneous soft tissue diffuses scalp-applied current in cadaver brains. a Schematic of the experimental arrangement for transcutaneous,
subcutaneous, and epidural stimulation. Example signal traces recorded in a coronal plane. Note the phase reversal of sinusoid voltage traces between thetwo sides. b Both transcutaneous and subcutaenous stimulation show intensity-independent linear (ohmic) properties (n = 81 in four different
arrangements in 10 cadavers, R = 0.92, P < 0.001 for subcutaneous, and n = 14 in 6 cadavers, R = 0.86, P < 0.001 for transcutaneous stimulation; raw data
and ﬁtted line are shown), which allows the calculation of voltage–current relationship. c, d Subcutaneous stimulation (c, R = 0.56, P < 0.001, n = 29 in 10
cadavers) elicited several-fold larger intracerebral gradients compared to transcutaneous stimulation (d, R = 0.8, P < 0.001, n = 16 in 6 cadavers).
Extrapolation from the measured data indicates that approximately 6 mA transcutaneous current can induce 1 mV/mm intracerebral electric ﬁeld (circle).
Raw data and ﬁtted lines are shown. e Ratios of induced intracerebral ﬁelds and stimulus amplitude in trancutaneous vs. subcutaneous (P < 0.001, n = 36 in
two different arrangements in 6 cadavers), and subcutaneous vs. epidural stimulation (P < 0.001, n = 60 in 3 cadavers). f 58 ± 7% of the applied current isshunted by skin and soft tissue and a further 16 ± 8% is attenuated by the serial resistance of the skull. g Effect of skull thickness on induced ﬁelds (n = 64
in 8 cadavers)
Fig. 4 Measuring induced intracerebral electric ﬁelds in human cadavers. a Photomicrograph of the custom-made multicontact electrodes used in the
cadaver experiments. b Photograph of the skull with drilled holes and inserted electrodes. A needle electrode in the sagittal plane on the forehead served as
reference for the recordings. Ag/AgCl electrodes, marked by blue and red circles for negative and positive polarity, respectively, were ﬁxed to the skull byconductive gel and secured by rubber-lined washers ﬁxed to the plexiglass frame by threaded rods. c Stereotaxic coordinates of the electrode shanks.
Numbers denote the number of recording sites for each electrode shank. Electrode tips (and adjacent sites) were positioned at the same depth to sample
distinct horizontal planes. The depth coverage of our electrodes was 3–7 cm (depending on the number of contact sites). d The effect of different
stimulation electrode conﬁgurations on the distribution of voltage gradients displayed on a single horizontal slice. Position of the stimulating electrodes
determines the location of maximal intracranial effect. Voltage gradients were calculated the same way as in Fig. 1b. e–g Effect of stimulus intensity,
frequency, and electrode size on intracerebral voltage gradients, respectively. Top three panels denote example gradient maps in the horizontal plane,
bottom graphs show population data. e Electric ﬁeld strength is a linear function of applied stimulus intensity (R = 0.52; P < 0.001; n = 48 gradient values in
two different arrangements in 4 cadavers). f Stimulus frequency between 5 and 1000 Hz has a minor effect on intracerebral gradients (one-way ANOVA; F(8, 891) = 0.0667, P = 0.99, mean ± 2 SD is shown, n = 900 gradient values in 5 cadavers). g Increasing electrode size increases the magnitude of electricﬁelds in a constant voltage mode, as the aggregate resistance decreases (n = 60 from 2 cadavers, P < 0.001 for all conditions)

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aLeft hemisphere

Right hemisphere

36
37

Trials

35

34

P3 P4

38

250 μV /7.5 mA

Left

+
0
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0

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* ***
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150
100
50

150
100
50
0 1.5 3 4.5 6 7.5
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0 1.5 3 4.5 6 7.5
TES intensity (mA)

Fig. 6 ISP stimulation of the scalp phasically modulates ongoing brain oscillations in human subjects. a Five consecutive example trials demonstrating alphaamplitude increase for high-intensity ISP stimulation. Alpha-band ﬁltered EEG signals are color coded based on the instantaneous ISP phase for better
visibility; blue and red colors denote stimulus trough (right-to-left currents) and peak (left-to-right currents), respectively. Gray sinusoids denote the ISP
stimulus epoch with an increasing–decreasing amplitude. b Phase modulation of the alpha amplitude by ISP stimulation for the entire session from thesame subject as shown in a, showing intensity-dependent alpha amplitude increase (mean across phases are tested in n = 45 trials against 0 mA condition,
P < 0.001 for 6 and 7.5 mA). Note the alternating phase modulation of the left and right hemisphere-derived EEG signals at 6 and 7.5 mA intensities. Colormaps show the phase-dependent median alpha amplitudes. c Population analysis for the left and right hemispheres, respectively, revealed an intensitydependent effect. Alpha amplitudes at stimulus peaks and troughs were generally unchanged for stimulus intensities below 4.5 mA. In response to anodal
currents in the same hemisphere, phasic modulation was signiﬁcant at 4.5, 6, and 7.5 mA. In response to cathodal stimulation in the contralateralhemisphere, signiﬁcant effects were observed only at 7.5 mA (right hemisphere) or 6 and 7.5 mA (left hemisphere)

that, similar to the rat, the highest electric ﬁelds occurred in the
neocortex near the stimulation electrodes (Fig. 4d; Supplementary
Figure 5). The generated gradient patterns were independent of
whether voltage or current mode was used for stimulation(Supplementary Figure 6a), and the relationship between appliedcurrent or voltage and the measured electric ﬁelds was linear(Fig. 4e; Pearson’s linear correlation; R = 0.52; P < 0.001; n = 48).
The frequency of stimulation had only a small effect on themagnitude of the induced ﬁelds (Fig. 4f; one-way ANOVA; P =0.99; F(8, 891) = 0.06; n = 900 trials from 5 cadavers). These
results demonstrate ohmic properties of the brain34, the surrounding skull, and soft tissue with negligible capacitive components35. When multiple electrodes were simultaneously
stimulated, the common conductive soft tissue summed the
independently applied dipoles, as shown by an equivalent circuit
diagram (Supplementary Figure 6b). As a result, simultaneously
applied multiple stimuli at different locations but with anincreasing phase separation gradually decreased the magnitude ofelectric ﬁelds (Supplementary Figure 6c). As expected, larger size
electrodes induced larger electric ﬁelds (Fig. 4g, case 1: 0.94 ±0.041, 1.25 ± 0.05, 2.84 ± 0.097 mV/mm, P < 0.001 for all comparison; and case 2: 0.23 ± 0.01, 0.32 ± 0.017, 0.43 ± 0.09 mV/mm, P < 0.001 for all comparison; mean ± SEM; n = 60 gradient8

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values)36. Unilateral placement of the electrodes restricted theinduced electric ﬁelds largely to the selected hemisphere (Fig. 4d;
Supplementary Figure 7; Etargeted = 0.9 ± 0.06 mV/mm and Enontargeted = 0.49 ± 0.03 mV/mm, mean ± SD; n = 30 cortical electricﬁeld values), similar to the results obtained in rats (Fig. 1b).
Figure 5 presents comparisons between transcutaneous (scalp),
subcutaneous (skull with scalp removed), and direct epidural
stimulation results (Supplementary Table 3). Transcutaneous
experiments used a limited set of recording electrodes to one
plane of the 3-D matrix (total of 28 contact sites on fourelectrodes in 6 cadavers) introduced via small individual incisionsof the otherwise intact scalp. The voltage–current relationshipremained linear for transcutaneous stimulation as well but theslopes were strongly reduced (Fig. 5b; Pearson’s linear correlation;
Rsubcutaneous = 0.92 and Rtranscutaneous = 0.86; P < 0.001 in bothcases; n = 14 and 81 for transcutaneous and subcutaneousconditions, respectively). Comparison of the current–electric ﬁeld
relationships indicated that approximately 2 mA subcutaneously
applied current was sufﬁcient to induce a 1-mV/mm ﬁeld
maximum (Fig. 5c; Pearson’s linear correlation; R = 0.56; P <0.001; n = 29). In contrast, the current vs. electric ﬁeld slope wasdecreased three-fold when current was applied to the scalp rather
than to the skull (Fig. 5c, d; Esubcutaneous(mV/mm) = 0.41 × I(mA)| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3+ 0.15; Etranscutaneous(mV/mm) = 0.13 × I(mA) + 0.04). Extrapolation of transcutaneous stimulation results suggested thatapproximately 6 mA current applied to the scalp would beneeded to reach 1 mV/mm voltage gradient in the living brain(Fig. 5d; Pearson’s linear correlation; R = 0.80; P < 0.001; n = 16).
Across all experiments in which scalp, skull, and epidural
stimulations were tested (n = 6), we could establish that 58 ± 7%of the current applied through the scalp diffused through the soft
tissue surrounding the head. Another 16 ± 8% of the current was
attenuated by the resistance of the skull, whereas current spread
effectively in the brain, including the meninges, vasculature,
ventricles, gray matter, and white matter (Fig. 5e, f; 0.12 (IQR =0.07–0.19) and 0.62 (IQR = 0.44–0.79) mV/mm/mA for transcutaneous vs. subcutaneous comparison; 0.61 (IQR = 0.49–0.80) and
0.93 (IQR = 0.67–1.23) mV/mm/mA for subcutaneous vs.

epidural comparison; paired t-test; P < 0.001 in both comparisons; n = 36 and 60 for the two comparisons, respectively),
supporting the view that the brain is an effective volume
conductor3. Skull thickness was a potential factor in attenuationof the current spread, explaining some of the variability acrosssubjects (Fig. 5g; Supplementary Figure 8; Pearson’s linear
correlation; P = 0.008; R2adjusted = 0.046; n = 128 electric ﬁeld
strength and skull thickness value pairs).
The experiments on the cadaver brains were performed from 3
to 8 days after death (Supplementary Table 1). Water density
measurements of brain specimens revealed that the postmortem
age had little effect on the hydration level of the brain(Supplementary Figure 9a). However, previous papers reported
that the resistivity is increased in the postmortem skull and soft
tissues37,38. To mimic the effects of postmortem changes on ourab 40Spectral amplitude (μV)

P3 lead - control

ISP stimulation
P3 lead - stimulated

100 μV

P3

Control
2 mA
4.5 mA
7 mA
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30

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Alpha band
spectral amplitude (μV)

100

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–0.2

Mean resultant length ofalpha band MI (z -score)**z -scored wavelet powereFrequency (Hz)Mean alpha amplitude (μV)

*
120

TES
Control - pre
2 mA
4.5 mA
7 mA
9 mA
Control - post

20

200

100 msd15Frequency (Hz)st
1 half of sessionsnd
2 half of sessions

200
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50
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0pre 2 4.5 7

9 0post0

TES intensity (mA)

4.5
1.5
3
6
TES intensity (mA)

7.5

Fig. 7 High intensity ISP stimulation of the scalp phasically modulates ongoing alpha waves in human subjects. a Blue, control EEG trace; red, EEG trace
during 7 mA stimulation (eyes closed in each condition). A 3-s magniﬁed segment of EEG trace at P3 lead is also shown. Note the absence of signal
saturation. The 1 Hz modulation of the baseline was removed. b Single session example of power spectra of EEG traces during increasing ISP intensities at 1
Hz. Power spectra were calculated in 10-s long windows, then averaged. c Quantiﬁcation of ISP stimulation-induced increase in alpha band power in asingle session. The control frequency band (120–140 Hz) showed no signiﬁcant change (n = 405–408 stimulus cycles; Palphaband = 0.37, 0.42, <0.001,<0.001; Pcontrol = 1.38, 1.38, 0.31, 0.62 for 2, 4.5, 7, 9 mA; all vs. 0 mA). d Single session example of alpha wave amplitudes as a function of the phase of 1
Hz ISP sinusoid stimulation. Asterisks denote phase bins signiﬁcantly different from the mean (one-way ANOVA; P < 0.05 at F(7, 3140) = 3.033 for 4.5
mA, F(7, 3136) = 6.96 for 7 mA, and F(7, 3160) = 14.37 for 9 mA; asterisks show signiﬁcant (P < 0.05) phase-intensity combinations of the post-hoc ttests against the phase bins of the control condition. e Single session example wavelet map (9 mA, 1 Hz ISP) shows ISP phase modulation of the alpha band
power. f Alpha band power modulation of wavelet decomposed EEG by 1 Hz ISP stimulation phase (n = 16, 10, 8, 18, 10, and 12 sessions; P = 0.98, 0.041,<0.001, 0.019, and 0.17 for ‘0pre’, 2, 4.5, 7, 9, and ‘0post’ intensities in 3 subjects; two-sample Kolmogorov–Smirnov test). g ISP stimulation-induced
increase of alpha power was stable throughout the recording epochs, as shown by the similar values during the ﬁrst and second halves of the stimulation
sessions (n = 23 trials from a single subject, P = 0.96, 0.79, 0.44, 0.44, 0.74, and 0.11 for 0, 1.5, 3, 4.5, 6, and 7.5 mA intensities, respectively)
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Eyes closed

Eyes open
Control

Control

2 mA

Control

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Control

2 mA

1 min

1 min

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L–R current
R–L current

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Abdominal
ISP

P4

Left hemisphere (P3)*** ***Δ Power (%)Δ Power (%)*** ***

6 mA

c

10

5

Control

P3

Right hemisphere (P4)

Left hemisphere (P3)15

P4Spectral power (μV2)

b

6 mA

Shuffled
ISP

ISP

P3

Control

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Frequency (Hz)

35

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Spectral power (μV2)

Right hemisphere (P4)
–10

–15

–20

2 mA 6 mA 2 mA 6 mA 6 mA
ISP ISP shuf shuf abdom

2 mA 6 mA 2 mA 6 mA 6 mA
ISP ISP shuf shuf abdom

600
400
200
0

5

10

15

20

25

30

Frequency (Hz)Fig. 8 Comparison of ISP and shufﬂed ISP stimulation on the EEG of human subjects. a Testing sequence of the experimental protocol. ISP stimulation used
the same arrangement as in Fig. 5. During shufﬂed ISP, adjacent stimulation electrodes were stimulated with opposite polarity. While shufﬂed ISP increases
local current ﬂow in the scalp, the alternating directions of the induced electric ﬁelds at the focus area are expected to cancel, ideally resulting in a zerocurrent in the brain. b Group results shown separately for the left and right hemispheres. Six mA current ISP stimulation increased alpha power in bothhemispheres. Shufﬂed ISP exerted only a unilateral and weaker effect. Intensities at 2 mA were ineffective. Abdominal stimulation (6 mA ISP protocol) did
not exert a signiﬁcant effect on alpha power (n = 809 epochs; power difference values for each conditions from 6 subjects, one-sample t-test with
Bonferroni correction). c Spectral power comparison between eyes open control and eyes closed control periods. Horizontal lines indicate signiﬁcantchanges from the eyes open condition (P < 0.05; n = 125 epochs for the eyes open, 144, 117, 126, 127, 148 epochs for the consecutive eyes closed, and 211
epochs for the abdominal ISP conditions; Mann–Whitney U-test with Bonferroni correction). Color coding of the conditions is the same as in a

cadaver experiments, we simulated the effects of changing theresistance of the soft tissue and examined the role of shunted
current by soft tissue (parallel resistance) and the current entering
through the skull (serial resistance). Our simulations showed thatthe lower resistances of the live soft tissues shunt the scalpapplied currents even more than in the cadavers (Supplementary
Figure 9b, c).
Overall, our measurements in human cadavers demonstratethat a signiﬁcant fraction of the scalp-applied current is lost due
to the shunting effects of the skin and soft tissue and the serial
resistance of the skull. Approximately three quarters of thecurrent was attenuated across the scalp and skull. These ﬁndings
were further supported by measuring the induced voltage ﬁelds
ﬁrst in vivo, followed by identical measurements up to 5
postmortem days in chronically implanted rats (n = 3; Supplementary Figure 9d).
Affecting human brain network activity by ISP. Our humancadaver and in vivo rat measurements indicated that approximately 6 mA currents applied to the scalp are needed in humans
to effectively alter brain networks. To test this prediction, a10

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circular array of 12 stimulation electrodes (six on each side;
Fig. 5a) was placed around the head and an ISP protocol wasapplied in 19 healthy human subjects. Each stimulation site
consisted of a 0.9% NaCl solution-soaked sponge square connected to 2 × 3 cm copper mesh. Scalp EEG was monitored by a
2-site montage (P3 and P4 against reference Pz). In each session,
1-min long control recordings were obtained before and after the
stimulation session (12 min). To avoid onset and offset effects,
ISP stimulation consisted of a train of 1-Hz sinusoids withincreasing and decreasing intensity (0, 1.5, 3, 4.5, 6, 7.5, 6, 4.5, 3,
1.5, 0 mA per cycle; Fig. 6a, b; Supplementary Figure 2b) for 12 s,
repeated 60 times for each subject. The low frequency (1 Hz)
stimulation allowed us to investigate the anodal–cathodal phasemodulation of the amplitude of the spontaneous EEG (represented mainly by the dominant alpha band activity) simultaneously in the left and right hemispheres. This approach exploits
the observation that TES can coherently affect the membranepotential (Vm) of many similarly oriented neurons (Fig. 2) andthe one order of magnitude lower stimulation frequency compared to the oscillation being observed reduces the possibility that
stimulation artifacts and their harmonics contaminate the| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3

results4. The residual ISP artifacts were removed by an ofﬂine
subtraction of the stimulation-triggered moving average (Supplementary Figure 1a). The artifact-removed signal preserved the
major features of the unstimulated control brain activity, as
demonstrated by the non-zero peaked cross-correlograms and theweakly correlated instantaneous frequencies in the alpha band
between the two EEG channels (Supplementary Figure 1b, c).
TES phase modulation of the amplitude of alpha waves became
visible by eye on the ﬁltered signal at high ISP intensities (6 and7.5 mA; Fig. 6a, b; paired t-test with Bonferroni correction; P <0.001 for 6 and 7.5 mA in both hemispheres; n = 45 trials). The
LFP modulation was present in both hemispheres and alternatedin phase, due to the shifting of the anodal–cathodal current
direction (compare blue and red epochs in Fig. 6a, b). For group
statistics, the mean alpha amplitudes near the stimulus peak(−135° to −45°) and near the stimulus trough (45° to 135°) were
measured separately at P3 and P4 at each current intensity.
Signiﬁcant modulation of the LFP amplitude by the TES phase
was observed at current intensities of 4.5, 6, and 7.5 mA at eachhemisphere when the preferred current direction was applied(Fig. 6c, paired t-test with Bonferroni correction; left hemisphereat TES trough: P = 0.006, <0.001 for 6 and 7.5 mA; left
hemisphere at TES peak: P = 0.01, <0.001, <0.001 for 4.5, 6,
7.5 mA; right hemisphere at TES peak: P < 0.001 at 7.5 mA; righthemisphere TES trough: 0.01, <0.001, <0.001 for 4.5, 6, 7.5 mA;
n = 1025 trials for all conditions from 18 subjects. All intensities
were tested against the 0 mA condition).
Stimulation intensity of 4.5 mA or higher induced various selfreported subjective effects, including tingling and slight burning
feeling of the skin in each subject. Subjective feeling of headmovement (dizziness) and horizontally oscillating light in thevisual ﬁeld at the frequency of the stimulation was often reported
even though the eyes were closed and testing was performed insemi-darkness. Feeling of a moving noise source in the horizontal
plane at 1 Hz was also reported. Nearly all subjects also reported a‘metallic taste’ in the mouth. All subjective effects were stronger atthe beginning of the stimulation session and attenuated (though
they did not disappear) by the end of the session. No poststimulation aftereffects were reported after any session. One
subject requested an early termination of the stimulation due todizziness and their data were excluded from the study.
In three of the subjects, we also used step currents of 1 Hz ISP
stimulation, instead of the intensity increasing–decreasing ramp.
The results of this experiment support those obtained by ramp
stimulation (Fig. 7a–f). Current intensity exceeding 4.5 mA
increased alpha band spectral amplitude (Fig. 7b–c; paired t-test
with Bonferroni correction; P < 0.001 for 7 and 9 mA vs. 0 mA; n= 405–408 power values for each stimulus intensity), and brought
about stimulus phase-dependent amplitude modulation of thealpha waves (Fig. 7d–f; two-sample Kolmogorov–Smirnov test; P< 0.001 and 0.019 for 7 and 9 mA vs. 0 mA; n = 16, 10, 8, 18, 10,
and 12 modulation vector length). This experiment also showedthat the subjective decrease of the perceived sensory effects overthe course of the experiments cannot be attributed to changes inalpha power since alpha power did not change between the ﬁrst
and second halves of the experiment (Fig. 7g; paired t-test; P =0.96, 0.79, 0.44, 0.44, 0.74, 0.11 for 0, 1.5, 3, 4.5, 6, 7.5 mA
intensities; n = 23 spectral amplitude pairs).
In a separate group of subjects (n = 6, Supplementary Table 4),
we examined whether the EEG changes could be accounted by
non-speciﬁc sensory stimulation or potential arousing effects of
TES. We used current steps (tACS) of 1 Hz ISP stimulation (2 or
6 mA for 5 min with 1-min rest periods; Fig. 8a). To test for theeffects of peripheral sensory stimulation, we interleaved the ISP
stimulation epochs with a similar protocol, where the adjacentelectrode pairs were stimulated with opposite polarity (‘shufﬂed
NATURE COMMUNICATIONS | (2018)9:483

ISP’ protocol—ﬁrst pair received an L–R direction current pulse,
second one received an R–L, third one an L–R, and so on; Fig. 8a).
Because of the alternating direction of the induced electric ﬁelds
with the shufﬂed ISP, the summed effect in neurons was expected
to be close to zero. The results with the regular ISP protocol
conﬁrmed the hemisphere-speciﬁc, stimulation phase-induced
modulation of alpha waves (Fig. 8b; Supplementary Figure 10;
one-sample t-test with Bonferroni correction; P < 0.005 for L–R
and R–L direction 6 mA ISP in both hemispheres; n = 809 power
difference values). In contrast, the shufﬂed ISP induced only
minor physiological effects on EEG activity (Fig. 8b; one-sample
t-test with Bonferroni correction; P < 0.005 for L–R 6 mA
shufﬂed ISP in left hemisphere; n = 809 power difference values).
Stimulus intensity at 2 mA failed to induce any detectable changes(ISP or shufﬂed ISP; one-sample t-test with Bonferroni correction; P = 0.36, 1.06, 2.22, and 1.45; n = 809 power difference
values). As an additional control for potential arousal effects, thesame subjects were also tested with the same ISP protocol butwith the stimulation electrodes placed on the abdominal wall(Fig. 8a). No hemisphere-speciﬁc, stimulation phase-induced
modulation of alpha waves was observed in the peripheral control
experiment (Fig. 8b; 6 mA abdominal; one-sample t-test with
Bonferroni correction; P = 0.41, 0.39, 1.68, and 1.18; n = 809
power difference values).
The self-reported subjective effects in this last experiment,using the ISP protocol with step increase of tACS, were similar to
those observed with the ramp stimulation (Supplementary
Table 4). In contrast, under the shufﬂed or the abdominal ISP
protocol no subject reported phosphenes, vestibular effects, or
metallic taste. On the other hand, each subject judged that the
feeling of skin ‘burning’ was stronger under the shufﬂed ISP
protocol. No post-stimulation aftereffects were reported by any
subject.
Finally, we asked whether the ISP or shufﬂed ISP stimulation
exerted a long-lasting blocking effect on occipital alpha oscillations. Alpha dominance of the EEG was present after eachstimulation period (eyes closed) and each period was signiﬁcantly
different from the control pre-stimulation eyes open condition(Fig. 8c; Mann–Whitney U-test with Bonferroni correction; P <0.05; n = 125–148 spectral power values for each of the 500
discrete frequencies). Some periods also showed signiﬁcant
increase in beta power (16–28 Hz) and decreased power in thedelta (<4 Hz) band compared to the eyes open condition. None ofthe post-stimulation periods were different from the prestimulation eyes closed control condition. Overall, these experiments showed that ISP stimulation at sufﬁcient stimulus intensity(4.5–6 mA) can bring about stimulus phase-dependent amplitude
modulation of the EEG, an effect which cannot be attributed toperipheral sensory or arousal effects.
Discussion
Using in vivo experiments in rodents and high-density intracerebral recordings in human cadaver brains, we found that scalpapplied currents were halved by the skin and subcutaneous soft
tissue and further attenuated by the skull by 16%. To reduce thescalp stimulation side effects and enable recording of brain
activity during stimulation, we designed an ISP stimulation
method that allowed injection of higher current intensities into
the brain compared to conventional tACS and transcranial direct
current stimulation (tDCS) approaches, while keeping the
cumulative charge density on the scalp surface and sensation
relatively low. The spatial speciﬁcity of the ISP method wasdemonstrated by hemisphere-speciﬁc activation of single neurons
in rats. Finally, we show that sufﬁciently strong induced ﬁelds can| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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directly and instantaneously affect brain networks in healthy
subjects.
Neuronal excitability is largely determined by ionic conductances brought about by neurotransmitter-induced postsynaptic potentials. However, neurons can also sense electricﬁelds directly1–3,28,39. In principle, the effectiveness of theexternally applied ﬁeld depends on a variety of factors including
neuronal density and geometry, alignment of dendrites and axons
relative to the induced ﬁeld, type and distribution of ion channels
in the neurons, degree of myelination, and density of glia, which
can affect the electric shunting effect of the extracellular space.
These factors can change between brain regions and might varyacross species and their potential signiﬁcance must be addressed
experimentally. The additive nature of the two polarizing
mechanisms (synapse-mediated and ephaptic) can be exploited toinﬂuence subthreshold membrane potentials and spiking1. Whena neuron is about to emit a spike, even a weak electric ﬁeld canbias spike threshold. In vitro experiments have shown that coupling an oscillatory ﬁeld to intracellularly generated oscillation
can be effective with as small as 0.2 mV/mm gradient40 and 1
mV/mm can induce measurable subthreshold Vm effects andaffected timing of action potentials39. Whether such weak stimuli,
applied acutely or chronically, can have beneﬁcial or deleterious
effects on brain function can only be determined by targetedrecordings and additional behavioral measures. To place such
weak forced ﬁelds into perspective, voltage gradients of >4 mV/
mm are brought about by hippocampal theta oscillations acrossthe CA1 pyramidal layer and >15 mV/mm ﬁelds are induced bysharp waves1. However, to exert a reproducible and instantaneous
impact on local networks by forced ﬁelds, the discharge of at leasta fraction of neurons with common targets should be temporally
coordinated. Our in vivo intracellular recordings have revealed
that electric ﬁelds >1 mV/mm were needed to exert measurable
effects on spikes and subthreshold Vm, but several times larger
currents were required to measurably affect the associated network rhythms1,4,6,8,12,41. This difference may be explained by thecompetition between the applied ﬁelds and the strong inﬂuence ofendogenous network patterns.
Given the complex paths of current spread in the brain42 andthe importance of neuronal geometry in sensing ﬁelds43, theabsence of an effect in any experiment cannot be taken as evidence for absence of effects on a few neurons. Yet, it is also
important to emphasize that the requirements of affecting spike
threshold of some neurons occasionally in wide areas of the brain
vs. consistently biasing activity of neuronal circuits are different.
We focused our efforts on establishing a current intensity level
required to reliably affect local networks in the intact brain. The
ﬁndings demonstrate that to consistently affect spikes and local
networks, sufﬁcient magnitude of charges has to be applied to thescalp to achieve >1 mV/mm voltage gradient44. In our rodent
experiments, more neurons were excited than suppressed byalternating current simulation, and the depolarizing ﬁelds weremore effective than hyperpolarizing ﬁelds. These ﬁndings illustrate that externally applied alternating ﬁelds bring about adominant excitatory gain. We emphasize though that our resultscannot rule out the potential impact of long-lasting (minutes tohours) weak electric ﬁelds on behavior by yet unidentiﬁedmechanisms45.
Although computational methods have become increasingly
sophisticated over the years17,28,35, experimental data are needed
to justify the modeling assumptions. Subdural measurements in
implanted patients46 can be useful but limited for estimating thecurrents necessary to bias spiking because they estimate ﬁelds
tangentially to the cortical surface, whereas the largest voltage
gradients are oriented orthogonal to the cortical surface44. Using
scalp, cranial, and epidural stimulation electrodes and multiple
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recording electrodes, we quantiﬁed the 3-D spread of electricﬁelds in both rodents and human cadavers. Our ﬁndings conﬁrm
the largely ohmic nature of current spread in the brain34, skull,and the surrounding soft tissue35. The scalp, subcutaneous tissue,
and muscles function as an effective shunt, resulting in at least
50% reduction of applied current intensity. The serial resistanceof the skull further reduces the current ﬂow by another 10–25%,
depending on the thickness of the skull47. Given the importanceof these attenuating factors, the amount of soft tissue, hair, andskull thickness should be taken into account in estimating theeffective current reaching the brain35, and variation of thesefactors may explain the large individual variability of the TESinduced effects14,46,48.
In TES practice, currents larger than 2 mA are avoided because
higher intensities are associated with skin sensation, phosphenes,
and other side effects7,9,49. Our intracellular and extracellularmeasurements in rats, along with previously published data4,6
and models17,41, suggest that 1 mV/mm is needed to affect neuronal ﬁring consistently and even larger strengths may be neededto phase-entrain brain rhythms to arbitrary stimulus frequencies.
In our cadaver experiments, we found that scalp-applied currents
should exceed 4–6 mA to achieve 1 mV/mm voltage gradient inbrain tissue but direct comparison with in vivo brain measurements are not available. Electric conductivity of the postmortem
tissue may change after death. However, our in vivo vs. postmortem comparison in rats demonstrated that in the intact head,actually larger currents are needed to achieve voltage gradientscomparable to those of the cadaver experiments, possibly due tothe stronger shunting effect of the better hydrated in vivo tissue38.
To reduce scalp sensation or other side effects31,41 and toincrease the direct effects of TES on brain activity7,9–14, novel
approaches are needed18. Our ISP method exploits the timeintegrating property of the neuronal membrane (i.e., the membrane time constant of neurons is ~10 ms), by applying brief andrapidly rotating ﬁelds. Using the ISP protocol, the strongest
transmembrane charge builds up where successively induced
electric ﬁelds (‘beams’) intersect, although in the human brain theanisotropy brought about the cortical gyri and ventricles should
be taken into consideration. The more ‘beams’ are used, thesmaller the adverse effects on other areas. Using just threerotating dipoles in rats, we demonstrated a proof of principle forthe focusing effect of ISP by conﬁning the ISP effect to largely one
hemisphere. In addition, when adjacent electrodes were stimulated with opposite polarity in human experiments (shufﬂed ISP),
no reliable brain responses were detected, even though skinsensation side effects increased. The latter effect may be explainedby the stronger current density induced by the opposite polarityof neighboring sites.
An experimental advantage of the ISP technique is that highfrequency pulses do not saturate recording ampliﬁers. This feature allowed us to measure the physiological effects of scalp stimulation in human subjects. Instead of focusing on brain
rhythm-entrainment effects15,50, in which residual artifacts arenotoriously difﬁcult to eliminate15,23, we examined how theamplitude of the spontaneous LFP was biased by the slowly
changing forced ﬁelds4. This method is similar to using tDCS at
multiple current levels, where the additive/subtractive effect of theapplied ﬁeld can be probed on the amplitude of native network
patterns4. We have veriﬁed the validity of this approach previously in rodents, using both LFP and unit ﬁring4. In support ofthe estimated voltage gradients from the cadaver experiments andthe ‘minimum’ ﬁelds (~1 mV/mm) in rodents to affect network
activity, we found that >4.5 mA currents were required to reliably
bias the amplitude of occipital alpha waves. While we designed
our experiments to maximize the stimulation effects on theparietal–occipital region where alpha waves are of largest| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3

amplitude, we cannot exclude the possibility that an improved
conﬁguration of stimulation sites could reduce the minimum
effective current somewhat further in future studies.
In our human experiments, we used six pairs of scalp stimulation electrodes, reducing the required local momentary current
by six-fold. Still, the effective intensities needed to directly bias
intracerebral neuronal activity induced adverse skin effects andvestibular reactions. An obvious next step in advancing the ISP
technique is to increase the number of intersecting dipoles generated by pairs of stimulating electrodes. For example, using amontage of 32 electrodes with highly conductive coupling to theskin, a large number of dipoles can be formed to create a circumscribed 3-D intersectional focus or target two or more brain
structures while reducing the locally applied currents, potentiallybelow the skin sensation threshold. Combining ‘ground truth’
measurements from the human cadaver brain with computationalmodels of the head can lead to a rationale design of focusedelectric activation of brain structures without adverse and perceivable peripheral effects.
In conclusion, we demonstrate that affecting neuronal circuits
directly and instantaneously in the human brain requires higher
intensity currents than used in conventional TES experiments.
Implicitly, our results also suggest that behavioral and cognitiveeffects reported in previous tACS studies have likely beenachieved by indirect mechanisms on brain activity, which needs
to be explored in detail. To achieve sufﬁcient magnitudes of
intracranial ﬁelds without direct peripheral side effects, novel
methods will be required.
Methods
Ethical permissions. The experiments were approved by the Institutional Animal
Care and Use Committee of New York University Medical Center IACUC (protocol number: 160926–01), the Ethical Committee for Animal Research (ethicalpermission numbers: XIV/471/2012 and XIV/218/2016), and the Ethical Committee for Human Research (ethical permission numbers: 98/2013 and 164/2014,
for the measurements on cadavers and healthy subjects, respectively) at the Albert
Szent-Györgyi Medical and Pharmaceutical Center of the University of Szeged inaccordance with European Union guidelines (2003/65/CE).

Subcutaneous tACS was performed in voltage controlled mode, generated by an
STG4008–16mA (Multi Channel Systems, Reutlingen). To monitor the appliedcurrent, a 100 Ω resistor was placed in series with the stimulating electrodes andthe voltage drop were measured across the resistor by an isolation ampliﬁer circuit
to preserve the ground-independent stimulation conﬁguration. The prefrontal
electrode was used as an anode and the lateral ones served as cathodes. Varying
frequencies (10, 20, 50, 100, 200, 500, 1000, and 2000 Hz) at 3 V were delivered
through the stimulating electrodes paired in various conﬁgurations. Recorded
signals (n = 30 channels) were ampliﬁed (10× gain) and stored after digitization at
20 kHz sampling rate per channel (RHD2000 Evaluation System, Intan
Technologies, Los Angeles).
Measuring the spatial selectivity of focused ISP in vivo. Two custom-designed(AutoCad, San Rafael, CA, USA) stimulation strips were 3-D printed (Form 1+,
Formlabs, Somerville, MA, USA) and glued bilaterally on the surfaces of thetemporal bones of the rats by cyano-acrylic glue (Loctite, Henkel). Each of the twosymmetric strips (width 13 mm, height 3.3 mm, and wall thickness 0.7 mm) consisted of ﬁve individual pockets which were separated from each other by 3.7, 2.2,
2.2, and 3.7 mm (Fig. 2b), and their medial surfaces were resembling the temporalbone curvature of a magnetic resonance imaging (MRI) data-based 3D model of a
rat skull. The middle pockets were positioned at 5.16 mm posterior from bregma.
The pockets were ﬁlled with conductive paste through ﬁlling holes left open at the
top (Super Visc, Brain Products, Germany) and then sealed with silicon. Craniotomies were drilled (2.2 mm diameter) and two silicon probes (Buzsaki32-H32;
NeuroNexus, Ann Arbor, MI, USA) were implanted at 5.16 mm posterior frombregma and 4 mm lateral of the midline, in the CA1 region of the hippocampus.
The hole around the probes was ﬁlled with non-conductive silicon (Dow Corning®,
Midland, MI, USA). Proper locations of the electrodes were conﬁrmed by thecharacteristic electrophysiological landmarks of the broadband signal at the pyramidal layer of CA1. ISP stimulation was performed in a voltage-controlled modeusing phototransistor-based custom-made electronics described below.
Comparing the effect of TES and ISP stimulation. To compare the effects of ISP
and DC stimulation in rats, the same surgery procedure was applied but thestimulation was performed in current-controlled mode (stimulus intensity 200 µA)
using the high-speed analog switch-based circuits described below. The recorded
signals (n = 64 channels) were ampliﬁed (400× gain) and stored after digitization at
20 kHz sampling rate per channel (KJE-1001, Amplipex, Szeged, Hungary). We
repeated the same measurements on one awake, freely moving animal4.

Experiments on rats. A total of 16 female, 3 male Long-Evans rats (350–450 g,
10–16 weeks old) and 8 male Wistar rats (250–450 g, 8–12 weeks old) were
implanted with custom-made recording and stimulating electrodes under urethane
anesthesia (1.3–1.5 g/kg, intraperitoneal injection) for the extracellular and thewhole-cell patch clamp recording experiments, respectively. Sample size for eachexperiment was estimated on the basis of anticipated inter-animal neurophysiological variability and the expected high success rate of the experiments in terms ofthe number of recorded neurons and data quality based on our previous studies.
No rats were excluded from the analysis. Each animal served as its own control, no
randomization or blinding was employed. Rats were kept on a regular 12 h–12 hlight–dark cycle and housed in pairs before implantation. No prior experimentation had been performed on these rats. After anesthesia induction, atropine (0.05
mg/kg, s.c.) was administered to reduce salivation, and the rectal temperature was
kept constant at 36–37 °C with a DC temperature controller (TMP-5b; Supertech,Pécs, Hungary). Stages of anesthesia were maintained by conﬁrming the lack ofvibrissae movements and nociceptive reﬂex. Skin of the head was shaved and theremaining fur was completely removed by using depilatory cream.

Comparing transcutaneous and subcutaneous stimulation. For transcutaneouselectrical stimulations, a pair of silicon single-pocket electrodes ﬁlled with conductive EEG gel was glued on both sides of the head of the rats, as described above(Fig. 1c). Small incision on the scalp was then made for craniotomy after subcutaneous lidocaine injections. Small craniotomy was drilled (1.2 mm diameter)
and a silicon probe (Buzsaki32-H32; NeuroNexus, Ann Arbor, MI, USA) was
inserted in the axis of the stimulating electrodes at 3 mm posterior from bregma
and 2 mm lateral of the midline, into the CA1 region of the hippocampus. A 50-µm
insulated wire electrode (California Fine Wire, Grover Beach, CA, USA) was
attached 1.2 mm away from the fourth shank serving as a reference electrode. The
hole around the probe was ﬁlled in with silicon (Dow Corning®, Midland, MI,USA).
After the transcutaneous stimulation, the silicon probe was removed and theskin was retracted and another set of silicon pocket electrodes were attached onto
the temporal bone, as described above. The silicon probe was inserted again atalmost the same location (2.8 mm posterior from bregma and 2 mm lateral of the
midline).
Varying frequencies (10, 100, and 1000 Hz) at varying amplitudes (10, 20, 50,
100, and 200 µA) were used for both settings in current-controlled mode
(STG4002; Multi Channel Systems, Reutlingen, Germany).
The recorded signals (n = 32 channels) were ampliﬁed (400× gain) and stored
after digitization at 20 kHz sampling rate per channel (KJE-1001, Amplipex,
Szeged, Hungary).

Recording intracerebral electric ﬁelds in anesthetized rats. To record thestimulus-induced intracerebral electric ﬁelds, the skin was retracted after a mediosagittal incision, and the bone surface was cleaned and dried as described
before51. Two strips of three custom-molded silicon pockets (2-by-2-by-1 mm, 4
mm2 surface area, 1.2 mm spacing) were glued onto the temporal bone bilaterally,
and a single pocket above the prefrontal bone (2 mm anterior from bregma in the
midline) using cyano-acrylic glue. Pockets were ﬁlled with conductive paste (Super
Visc, Brain Products, Germany) and then sealed with silicon. An Ag/AgCl reference electrode was placed in the subcutaneous space behind the neck. Thirty holes(0.5 mm diameter) were drilled in the skull and a custom-made 6 × 5 recording
electrode matrix was inserted into the brain 3 mm deep below the brain surface(Fig. 1a). The spacing between the individual electrodes was 2, 1.7, 2.2, 1.7, and 2
mm in the x axis and 2 mm in the y axis. Each recording electrode was made by apolyurethane-insulated, copper–nickel wire (50 µm diameter) inserted into a supporting polyimide tube (70 µm interior diameter, 86 µm outside diameter).

Measuring the effect of postmortem age. Rats were implanted with a pair ofsilicon pocket electrodes as described above. Twelve holes (0.5 mm diameter) were
drilled in the skull and a custom-made 6 × 2 recording electrode matrix wasinserted into the brain. The spacing between the individual electrodes was 2, 1.7,
2.2, 1.7, and 2 mm in the x axis and 2 mm in the y axis. The electrode matrix wasinserted at 3 mm depth in the brain and the craniotomies were ﬁlled with silicon(Dow Corning®, Midland, MI, USA). Once the silicon dried, the whole skull wascovered by dental cement (Duracryl™ Plus, Spofa Dental, Jičín, Czechia) and theskin was closed by sutures to restore its conductive integrity. Subcutaneous tACS
was performed in voltage-controlled mode using various stimulation parameters as
described above (STG4002; Multi Channel Systems, Reutlingen, Germany).
After the in vivo measurement, the rats were euthanized by sodium
pentobarbital (150 mg/kg, intraperitoneal injection). The corpses were kept at 4 °C
after death in plastic bags to prevent desiccation. Subcutaneous tACS andrecordings were repeated on postmortem day 1, 2, 3, 4, and 5. In one case, the

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implantation was done 5 days after the euthanasia, which resulted in qualitatively
and quantitatively similar results (data not shown).
In vivo whole-cell patch clamp recordings. A pair of silicon pocket electrodes
ﬁlled with conductive EEG gel were glued bilaterally on the skin or on the temporalbone of rats for transcutaneous and subcutaneous stimulation, respectively, similarly to extracellular recording experiments (Fig. 1c). A small craniotomy (~2 mmdiameter) was made 5.0 mm posterior from the bregma, 4.0 mm lateral of the
midline. Patch-pipettes were made from borosilicate glass capillaries (GC150TF-10;
Harvard Apparatus, Holliston, MA, USA) and their tip resistance were 5–7 MΩwhen ﬁlled with an intracellular solution: (in mM) 135 K-gluconate, 10 HEPES, 10
Na2-phosphocreatine, 4 KCl, 4 ATP-Mg, and 0.3 GTP-Na3 (pH = 7.25, 290mOsm). Liquid junction potential calculated as +18.6 mV was ofﬂinecompensated. The patch-pipettes attached to a ﬁne stepper motor manipulator
were lowered perpendicularly and blind in vivo whole-cell recordings from corticalneurons (0.5–1.3 mm from the pia) were obtained as previously described52.
Recordings were performed using an EPC10 USB ampliﬁer (HEKA Elektronik,
Lambrecht/Pfalz, Germany) with PATCHMASTER software (ver. 2.901 HEKA
Elektronik). Signals were ﬁltered at 3 kHz and digitized at 20 kHz. The pipettecapacitance, membrane capacitance, and series resistance were compensated. If
series resistance varied more than 20% or increased above 50 MΩ, the data werediscarded. Isolated direct constant current stimuli (~ ±800 µA) were delivered via a
multifunction stimulator (STG4002; Multi Channel Systems, Reutlingen, Germany). After the whole-cell transmembrane potential recordings, the recorded
neurons were detached from the pipette with retraction and positive pressures.
After retraction, the artifacts of the same set of electrical stimuli applied during thewhole-cell recordings were recorded extracellularly. The recorded artifacts weresubtracted from intracellularly recorded potentials to recover the true transmembrane potentials25. Finally, a 4-shank 32-channel silicon probe (Buzsaki32-H32;
NeuroNexus, Ann Arbor, MI, USA) was inserted in the vicinity of the recorded
neuron to record extracellular electrical gradients in response to the same stimulation as during intracellular recordings. The extracellular recordings were performed at 20 kHz sampling rate using a KJE-1001 ampliﬁer (Amplipex, Szeged,
Hungary).
Measurements on human cadavers. Recordings were performed at the Department of Pathology, Faculty of Medicine, University of Szeged. Medical history of
cadavers was consulted in advance and only those with no known brain disorderwere selected for measurements. The corpses were kept at 4 °C after death in plasticbags until autopsy to prevent desiccation. The autopsy theater temperature was 22 °C. The routine medical autopsy procedure was done on the same day as experimental measurements. The sample size required (number of cadavers andrecording sessions) was extrapolated from the results of animal studies. There wasno blinding or randomization employed. No cadaver was excluded from theanalysis.
Recording tACS-induced intracerebral electric ﬁelds. The scalp was cut along
the coronal plane connecting the mastoids. The anterior and posterior halves of thescalp were retracted forward and backward, respectively, until the supraorbital
ridge and the occipital protuberance was revealed. The temporal muscles and soft
tissue were also removed. After the skull was cleaned, the head was ﬁxed in acustom-made acrylic glass frame (Supplementary Figure 3). The top of the skull
was pushed against the acrylic frame as close as possible. Four stainless-steel screw
bars (6 mm diameter, 10 cm length) held the head steady on each side. One end of
the screw bar was attached to the tower by a hexagonal nut, the other end held arubber ring (3 cm diameter) against the skull. Once the head was positioned, thepositions of the 36 penetration holes were marked by an ink-ﬁlled needle throughthe pre-made hole-matrix of the plexiglass back panel. The frame was removed andholes were drilled (1.2 mm diameter) and rinsed by physiologic saline. The frame
was placed back to its original position, and the head was repositioned by the screw
bars. Four or seven pairs of stimulation electrodes (Ag/AgCl EEG electrodes, 10
mm diameter, Ambu, Copenhagen, Denmark) were placed between the rubber
rings and the skull surface with conductive paste (Ten20, D.O. Weaver, Aurora,
CO, USA). Custom-made multiple-site electrodes were prepared as follows: three
to seven holes were drilled on the outer surface of a translucent polyimide tube
(724 μm internal diameter, 775 μm outside diameter). Three to seven 127-μm
diameter, polyurethane-insulated, copper–nickel wires were threaded into thepolyimide tube through these holes space 1 cm from each other. The wires weresecured by a drop of cyano-acrylic glue at the side-holes of the polyimide tube andthe other end of it was soldered to a connector socket. The tubes were backﬁlledwith epoxy glue to increase stiffness. Once the epoxy dried, the wires were cut backat the surface of the polyimide tube, and the tip of the tube was sharpened.
Impedances of the contact sites varied between 50 and 300 kΩ at 1 kHz. Electrodes
were inserted into the brain through the previously drilled skull holes and thematching plexiglass matrix while rotating continuously, to preserve parallelalignment.
A needle was inserted through the skull above the prefrontal cortex and served
as reference electrode. Physiologic saline solution (2–5 ml) was injected through thesame hole to reﬁll the cerebrospinal ﬂuid lost during the drilling procedure.
Recording electrodes made a watertight seal in the skull holes, thus further leakage
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NATURE COMMUNICATIONS | (2018)9:483

was not signiﬁcant. The chest wall was used as grounding. Subcutaneous(electrodes placed on the skull surface) alternating and direct current stimulation
was performed using stimulation signals generated by either an STG 4008–16 mA(Multi Channel Systems, Reutlingen) or an NI 6343 board (National Instruments)
with precision isolation ampliﬁer circuits. The ﬂoating (cadaver) side of theisolation ampliﬁers were powered using two 9-V batteries for each stimulator pair.
To monitor the applied current, a 100-Ω resistor was placed in series with thestimulating electrodes and the voltage drop across the resistor was measured by anisolation ampliﬁer circuit. The stimulating electrodes of the two sides were paired
using different parallel or diagonal arrangements. Sinusoid stimuli with varying
intensities (1, 2, 3, 4, 5, and 6 V) at 10 Hz and varying frequencies (5, 20, 50, 100,
200, 500, 1000, and 2000 Hz) at 5 V were used for at least 200 cycles, each. To
mimic the effect of increasing electrode sizes, multiple stimulating electrodes were
coupled together. In some cases, an additional stimulation electrode was placed in
the midline of the forehead, and used against a selected lateral electrode to achievea fronto-lateral stimulation conﬁguration. The recorded signals (n = 198 channels)
were ampliﬁed (10× gain) and stored after digitization at 1.6 kHz sampling rate perchannel by a custom-designed circuit (for DC coupled recordings) or at 20 kHz
sampling rate per channel by another custom-designed recoding system based on
the RHD2000 Evaluation System (for AC coupled recordings, 0.1–6 kHz
bandwidth, Intan Technologies, Los Angeles, CA, USA).
Recording tDCS-induced intracerebral electric ﬁelds. Instead of the abovementioned electrode matrix, six custom-made single contact Ag/AgCl electrodes
were introduced in the fourth coronal plane in 2 cadavers. A silver wire (400 µm
diameter) was inserted into a translucent polyimide tube. The contact sites werecleaned by scratching with a razor blade and then immersed into NaOCl (42 g/l)
for 16 h. The same recording system was used to acquire signals (10× gain, 1.6 kHz
sampling rate per channel). All stimulating electrode pairs were active simultaneously and anodal or cathodal stimulation was applied for 50–50 s (5 V intensity).
Measuring the shunting effect of the skin and skull. Instead of retracting theskin, four or six 5-mm long incisions (15 mm apart from each other) were made on
both sides of the sagittal suture in the coronal plane, connecting one mastoid withthe other. To prevent soft tissue damage, the drilling head was used through ametal tube (1.3 mm diameter) and four holes were made. Then stimulation electrodes (n = 4, Ag/AgCl, Ambu, Copenhagen, Denmark) were attached to the skinby conductive paste (Ten20, D.O. Weaver, Aurora, CO, USA). Four or six custommade 7 contact site recording electrodes were inserted into the brain, transcutaneous alternating current stimulation was performed, as described above. The
recorded signals (n = 28 or 42 channels) were ampliﬁed (10× gain) and stored after
digitization at 15 kHz sampling rate per channel (RHD2000 Evaluation System,
Intan Technologies, Los Angeles, CA, USA). After the skin measurements, the skinincisions were carefully connected and the scalp was removed while the recording
electrodes were kept in place. The stimulating electrodes were attached to the skull
surface and the same stimulation protocol was applied. In separate experiments, tocompare the effect of subcutaneous stimulation to intracranial stimulations, insome cases additional stimulating electrodes were placed intracranially, in between
the subcutaneous electrodes as follows: the additional skull holes were drilled withincrementally increasing (2, 4 and 8 mm) drill-bit sizes, and externally threaded,hollow plastic dowels (15 mm long, 8 mm diameter, Hettich Furntech, Germany)
were introduced in the holes to form an electrical isolation toward the skull. Sponge
electrodes with the encapsulated Ag/AgCl plates, soaked in physiologic saline, wereglued to the tip of screws, and introduced into the plastic dowels to touch the brain
surface.
Registering the anthropometric data of the cadavers. At the end of the measurements, the cranium was opened with an oscillating saw in the line of thestimulating electrodes. After removing the skull cap, the brain was also removed.
Anthropometric data of the skull was measured (circumference, sagittal, horizontal,
vertical distance, and skull thickness below the stimulating electrodes). After thebrain was examined by the pathologist, a 5-g piece of the occipital lobe wasremoved to measure the water content of the brain tissue by desiccation. As
reference, hydration value of living tissue was taken from reference53.
Measurements on human subjects. Human transcutaneous stimulation andrecording experiments were performed on healthy subjects (all males, age = 21–66
years). Subjects with short hair were preferably selected, thus including only males
was incidental. All subjects gave their informed consent to the experiments. Eachsubject served as his/her own control; no randomization or blinding was used.
Before performing the ISP stimulation protocol, each subject was brieﬂy
exposed to a few seconds of 1 Hz constant current stimulation with increasing
intensities (1, 2, 4, and 8 mA) to familiarize them with the expectable subjective
experience during the ISP protocol, and to test if any adverse effects are present.
The intensity was increased to the next level only if the previous intensity wasreported as being well tolerable. In addition to the well documented tingling,burning feeling of the skin and perception of phosphenes11, stimulus intensities
above 4.5 mA stimulation induced feeling of horizontal head-movements andhorizontal oscillation of the visual and auditory ﬁelds at the frequency of the| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3

stimulation. All subjective effects were stronger at the beginning of the stimulation
and attenuated during the course of stimulation. None of these are considered
Serious Adverse Effect or Event by the US Food and Drug Administration54. No
aftereffects were reported after any session. Phosphenes were likely induced bycurrent spread through the orbits, whereas vestibular and auditory effects werelikely due to the spread of currents through the ear canals.
Considerations of TES effects on human subjects. There are no accepted
guidelines about the current limit of either tDCS or tACS24,49,54. The main reason
for this is the lack of reliable information about the induced ﬁelds in the humanbrain, and this is what we supply in our cadaver studies. Most TES studies use <2
mA, mainly because this is the threshold where peripheral sensation and phosphenes are typically detected. Other related measures include (1) current density(in A/m2) at the electrode calculated by taking the applied current to a given
electrode and dividing by electrode area and (2) stimulation charge (in Coulombs,
C) determined by multiplying current by duration. Since the adverse and risk
effects of stimulation are related to current density and duration of stimulation (i.e.,
the total charge or ‘dose’), 1 mA for 10 min, 2 mA for 5 min, and 10 mA for 1 min
are considered equivalent from the perspective of charge54, yet these three categories may not be equivalent for subjective side effects or instantaneous direct brain
effects. Direct stimulation of the brain via subdural electrodes using 1 ms pulses of
5 mA intensity for several seconds considered to be safe55. The brain-penetrating
currents used in our studies remained well below these widely accepted values. Oneof 19 subjects in the ramp stimulation experiments (Fig. 5) requested to terminate
ISP stimulation because of feeling dizzy. For the experiments shown in Fig. 6, we
recruited 7 subjects (3 subjects overlapped with the experiments shown in Fig. 5
but were tested several weeks apart). In one of them, the instability of the electrodes
was only discovered after the experiments and the results from this subject could
not be analyzed due to excessive artifacts.
Stimulation methods. Stimulating sponge electrodes for ISP were prepared from a
2 × 3 × 1.5 cm sponge glued to a 2 × 3 cm copper mesh, and glued to a rubber
washer with the sponges inside, keeping approximately 2.5 cm distance between
sponges. The rubber washer with the 12 electrodes was soaked in 0.9% saline
solution and tightened gently around the head. Conductivity was further improvedby putting electrode gel (SuperVisc, EasyCap GmBH, Germany) between the wetsponges and the skin. For abdominal ISP stimulation, the same sponge electrodes
were placed around the trunk.
Modeling of current-induced ﬁelds. To model the effect of soft-tissue resistivity
on intracerebral electric ﬁelds, a ﬁnite element method model was constructed, andthe theoretical values of the electric ﬁeld inside the brain were calculated for
different conditions using Comsol Multiphysics (Comsol, Burlington, MA, USA).
Concentric spheres simulated the scalp, skull, cerebrospinal ﬂuid, and brain(Supplementary Figure 9b). Conductivity values were set to 0.465, 0.015, 1.65, 0.3 s/m, respectively17,35. The dimensions of each layer were set to match one of the
cadaver’s anthropometric data (8, 5, and 2 mm thicknesses for the skin, skull, andcerebrospinal ﬂuid, respectively, while the brain diameter was set to 142.6 mm).
Two virtual stimulating electrodes attached to each side of the head were modeled
as conductors of 1 cm2 with the conductivity of copper (5×107 s/m). Induced
electric ﬁelds were calculated inside the model brain with a virtual 5 × 6 electrodes
array, mimicking the experimental setup. Maxwell’s equations were solved within
an adaptive mesh of 366619 elements, using a linear solver and a relative toleranceof 1e−6. The effect of skin and soft tissue resistivity change on electric ﬁelds was
similar to those reported in earlier publications investigating postmortem resistance changes of soft tissues and muscle37,38.
We used a leaky-integrate and ﬁre neuron model to visualize the ISP principle.
Extracellular electric ﬁelds were derived from in vivo tACS measurements using 1
kHz sinusoid stimuli in the same arrangement as shown on Fig. 1a, but using
epidural stimulation with screw electrodes. The directionless electric ﬁeld
intensities (35 mV/mm peak intensity) at each point were converted to
intracellularly injected current values by multiplying with a factor (4.5 nA/mV/mm) to mimic transmembrane currents. A dimensionless leaky-integrate and ﬁreneuron model was established in Matlab based on ref. 31. Parameters were set as
the following: temporal constant of the membrane = 10 ms; resting membranepotential = −70 mV; membrane resistance = 1 MΩ; spiking threshold = −54 mV;
spike peak potential = 20 mV; repolarization level = −80 mV. Extracellular electricﬁeld duration = 0.5 ms. The effects of three different magnitude current injections
on the ﬁring rate were demonstrated by the leaky-integrate and ﬁre neuron model
is illustrated in Fig. 2a.
Electronic circuit of ISP stimulation. For the ISP stimulation approach, both
positive and negative leads of the stimulus generators were connected to 12–12
TLP52-4 phototransistors (Toshiba, Japan). Bidirectional, ground-independent
conductivity was achieved the following way. Two phototransistors were seriallycoupled through their emitter and collector, and the input signal from the waveform generator was fed into both the emitter and the collector end of the transistor
doublet, through two Schottky-diodes, which allowed current ﬂow only to theappropriate member of the doublet, depending on the polarity of the signal. The
NATURE COMMUNICATIONS | (2018)9:483

common segment of diode, the doublet, was connected to a stimulation electrodeon the head. The same circuit was constructed for the other pole of the signal aswell. Common driver signal to the infrared emitting diode sides opened all fourtransistors, but due to the diodes two of them were always ﬂoating, while the othertwo closed the circuit through the head (Supplementary Figure 1a). Six such circuits were used for the six electrode pairs, forming six quadruplets (blocks) oftransistors. In rats, only three pairs were used. Blocks were activated in a pseudorandom order by transistor–transistor logic (TTL) pulses generated by a
CD74HC4017 Decade counter (Texas Instruments, USA), driven by a 100-kHz
TTL generator (ADG3051C, Tektronix, USA).
In sessions employing variable ISP intensities in human subjects, thephototransistors were replaced with ADG412 high-speed analog switches (Analog
Devices, Norwood, MA, USA) and the control TTL signals were generated by a
PIC18F4525 (Microchip, Chandler, AZ, USA) microcontroller and isolated by
ADuM1400 (Analog Devices) digital isolators. This circuit allowed the unrestricted
ﬂexible assignment of stimulus polarities to the electrodes. This latter circuitry was
also used for the experiments on rats when we compared the spatial effect of ISP
and TES pulses.
EEG recording during ISP stimulation. EEG scalp recordings were performed by a
16-channel V-Amp ampliﬁer and ActiCap BP active electrodes (Brain Products
GmBH, Germany). Impedances were measured online and adjusted to remain
below 20 kΩ by applying electrode gel. Electrodes were placed according to theInternational 10/20 electrode scheme. The broad dynamic range of the active
electrodes, and their buffering capacity allowed the low-noise transmission of EEG
signals and stimulus artifacts without on-head ampliﬁcation. To prevent thesaturation of the ampliﬁer, the output range of the active electrodes was matched tothe input range of the EEG ampliﬁer through custom-made voltage dividers.
Data processing. The recorded data were analyzed by custom-written scripts in
MATLAB (MathWorks, USA). Single unit analyses were performed on the timeseries of semi-automatically clustered and manually reﬁned unit clusters of
extracellular action potential waveforms, as described earlier (PHY56,57). Only wellisolated single units were used in the analysis. Units were categorized as putative
pyramidal cells or interneurons based on their physiological properties.
To measure the electric ﬁeld in cadavers and rats, 500 sinus cycles wereaveraged for each condition and then the peak-to-peak amplitude was measured
for each channel. The ﬁrst spatial derivative of these voltage signals was calculated .
In the silicon probe experiments, we measured the impedance of all contact sites
at 10, 100, and 1000 Hz (Intan recording software, Intan Technologies, Los
Angeles) and excluded those channels whose impedance values were higher than 2
MOhm: 500 sinus cycles were averaged for each condition and then the peak-topeak amplitude was measured for each channel and a mean shank voltage wascomputed. Finally, we calculated the ﬁrst spatial derivative of these potential values.
Statistical tests. Unless otherwise noted Student’s paired t-test with Bonferroni
correction was used for pairwise data comparison, Pearson’s linear correlation was
calculated for correlation analyses and mean ± SEM values are displayed with thefull data sets superimposed. Data with non-normal distribution are reported as
median and interquartile range (IQR). Boxplots with whiskers denote the medians,
interquartile ranges, and full ranges. For data sets with non-normal distribution,
non-parametric tests were used instead. Welch’s correction was applied when
variances were not equal. To preserve visibility of the ﬁgure panels, signiﬁcancelevels of <0.05, <0.01, and <0.005 are marked by one, two, or three asterisks,
respectively. For simplicity, P values smaller than 0.001 are reported as <0.001
instead of stating the absolute value. The detailed conditions, numeric results, andeffect sizes of the statistical comparisons are listed in the Supplementary Results.
Analysis of whole-cell patch clamp recordings. To remove the stimulation
artifacts, after retracting the patch-pipettes, the artifacts of the same set of electricalstimuli applied during the whole-cell recordings were recorded extracellularly. The
recorded artifacts were subtracted from the intracellularly recorded potentials torecover the true transmembrane potentials25. Power spectra of the stimulated andcontrol epochs were calculated on a trial-by-trial basis, using fast Fourier transform, before averaging. Spectra were whitened by the 1/f method.
Noise removal and analysis of EEG activity. The stimulus artifact was removed
ofﬂine by subtracting a triggered moving average (t = 10 epochs), followed bytriple-sweeps of 100th order zero phase-lag high-pass ﬁnite impulse response ﬁlter(f = 2 Hz) in MATLAB.
For analyses performed on the time domain (e.g., alpha amplitude) the artifactfree signal was ﬁltered in the alpha band with a zero phase-lag fourth-order
Butterworth ﬁlter. Instantaneous alpha amplitudes were determined by calculating
the magnitude of the Hilbert-transformed ﬁltered signal, and binned based on thecorresponding ISP amplitude and phase. Binned values were averaged acrossepochs. To estimate the amplitude of the remaining electrical noise time locked tothe epochs, signal was ﬁrst averaged across epochs, and then Hilbert transformed.
This approach preserved time-locked features. For frequency domain analyses,
spectral amplitudes were calculated using fast-Fourier transformation, and| DOI: 10.1038/s41467-018-02928-3 | www.nature.com/naturecommunications

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NATURE COMMUNICATIONS | DOI: 10.1038/s41467-018-02928-3

smoothed using a moving average ﬁlter (width = 2 Hz). 120–140 Hz was chosen asa control frequency band, as this range does not represent measurable physiologicaloscillatory signals on the scalp but would still mirror the presence of broadband
electrical artifacts. For time-resolved spectral analysis, spectra were calculated using
a multitaper fast Fourier transform on 1-s long consecutive segments. Spectra werewhitened by multiplying each frequency by the frequency value (1/f method).
Frequency-amplitude and phase-amplitude analysis of EEG. We employed twocomplementary analyses to assess the modulation of EEG amplitude by the phaseof the sinusoidal ISP stimulation current. Analyses were performed on 1-min-longconsecutive epochs, and the epoch results were pooled. First, we applied thecomplex wavelet transform using Morlet mother wavelets to calculate the amplitude and phase for a wide range of EEG frequencies. Wavelet amplitudes were
calculated from 1 to 30 Hz at 59 levels from the artifact-free EEG and wavelet phasefor 21 levels from 0.5 to 5 Hz at 15 levels from either the original EEG or asynthetic signal constructed from the stimulation pulses. Phase–amplitude crossfrequency coupling was quantiﬁed using a modulation index (MI58). To quantify
frequency–amplitude modulation, 2-D comodulograms were constructed with theMI values for every phase–amplitude frequency pair and the maximal MI in theband of interest was detected59. For phase–amplitude modulation, phase timeseries were binned into phase intervals and the mean wavelet amplitude was calculated for each of them and z-scored. Phase time-series were binned into phaseintervals and the mean wavelet amplitude was calculated for each of them59.
For a complementary phase–amplitude analysis performed on the time domain,
the estimated peak-to-peak amplitude values of the individual alpha waves werebinned based on the actual stimulus phase, and alpha amplitude values during thestimulus peak and trough bins (45, 90, 135, 225, 270, 315°) were compared to thealpha amplitude values present at the transitional phase (0° and 180°) bins using at-test.
Data availability. The data sets generated and analyzed during the current study
are available upon reasonable request from the corresponding authors for further
analyses.

Received: 1 December 2017 Accepted: 9 January 2018

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Acknowledgements
This work was supported by EU-FP7-ERC-2013-Starting grant (No. 337075), the‘Momentum’ Program of the Hungarian Academy of Sciences (LP2013-62), GINOP2.3.2-15-2016-00018, US National Institute of Health grants (NS90583, MH107396,
NS099705), the Mathers Foundation, EMBO (ALTF 147-2015), the Uehara Memorial
Foundation, EFOP-3.6.1-16-2016- 00008, EU H2020 ERC Marie Skłodowska-Curie grant(No. 734718 – “CoBeN”), and the National Brain Research Program 2 (No. 2017-1.2.1-

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NKP-2017-00002). We thank Brendon Watson and Anli Liu for comments on the
manuscript; Jennifer Gelinas, Dion Khodagholy, Péter Vörös, Máté Kozák, Árpád Kormányos, and Bereniké Rokszin for assistance with aspects of the experiments and Péter
Liszli and Zoltán Gingl for technical advice on electronics.

Author contributions
A.B. and G.B. conceived the project. M.V. performed the in vivo rat experiments. M.V.
performed the cadaver experiments assisted by A.B., K.B., T.Z., and B.I. A.B., G.K., Z.T.
K., and M.V. did the human recording experiments and analyzed rat and human neuraldata. A.O. did the modeling of the human head. A.F.R. performed phase-amplitude
analysis. Y.T. performed intracellular recordings on rats. A.B., G.B., and M.V. wrote thepaper. A.B., G.B., and M.V. served as ﬁrst subjects of ISP stimulation to gauge potential
adverse effects.

Additional information
Supplementary Information accompanies this paper at https://doi.org/10.1038/s41467018-02928-3.
Competing interests: A.B. is the founder and owner of Amplipex and Evobrain LLCs,
which manufacture biosignal ampliﬁers and stimulator devices. A patent application
about the ISP method has been ﬁled by A.B. and G.B. The remaining authors declare nocompeting ﬁnancial interests.
Reprints and permission information is available online at http://npg.nature.com/
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