1.

Introduction

Transcranial current stimulation (tCS) is a noninvasive form of brain stimulation that applies weak direct electrical currents to the brain through electrodes placed on the scalp. Different types of stimulation are available based on how the current delivered to the brain is modulated.¹ The most common approach is to hold the current constant during the stimulation period, this is known as transcranial direct current stimulation (tDCS). This method is commonly referred to as tDCS, but in order to avoid confusion with the optical method, we here use DC-tCS. DC-tCS induces bidirectional, polarity-dependent changes in cerebral excitability of humans^2,3 and has been shown to modulate cognition in both healthy subjects (improvement of cognitive function) and in patients (reversing the effects of cognitive deficits after a stroke).^4,5 Positive effects have also been shown in treatment of depression, chronic and acute pain, Parkinson’s disease, focal epilepsy, and for improving recovery after stroke.^6–9 A therapeutic application of this technique is appealing due to its relative low cost, to its demonstrated safety,¹⁰ and as a substitute for pharmacotherapy, especially for patients with poor drug tolerability.⁸

In order to define protocols of application and to optimize the individual dosage of stimulation, the effects of the stimulation on the cerebral activity must be monitored in real time.^11,12 Availability of a concurrent read-out of the stimulation effects could improve its effectiveness and, ultimately, a personalized stimulation driven by the neuromonitor feedback can be imagined.¹³ For this purpose, due to the complex nature of cerebral activity, multimodal monitoring is preferable,^14,15 such that different monitors integrated in a single set-up can allow to record the effects at different levels, from neuronal activity, to hemodynamics to systemic.¹⁶ The neuronal activity is evaluated by electroencephalography (EEG), which measures the synchronous activation of a large population of pyramidal neurons oriented perpendicularly to the scalp. On the other hand, cerebral hemodynamics can be used as a surrogate measure of the cerebral activation assuming that cerebral blood flow (CBF) increases in a region of the brain, where neurons and synapses are activated in order to meet the demand for more energy and that the amount of local increase of blood flow and the changes in oxygen saturation are related to the cellular energy consumption.¹⁷

Even if different options are available as neuromonitors of hemodynamics, functional near-infrared spectroscopy (fNIRS) appears to be the best suited for this application.¹⁸ The most accessible and commonly employed fNIRS technology uses continuous-wave sources (CW-fNIRS), where constant intensity light sources are employed to measure oxy- (${\mathrm{HbO}}_2$) and deoxyhemoglobin (Hb) concentration changes as surrogate measures of CBF. fNIRS has advantages that are fundamental for this application. In particular, it allows for continuous noninvasive measurements, does not require the immobility of the subject, and is not necessary to change the environment preferable for the stimulation to integrate CW-fNIRS in the protocol.¹⁸ It has indeed already been used as a monitor during DC-tCS, both in animals¹⁹ and on humans at rest^20–22 and during tasks.^22–24 It has also been used together with EEG to monitor DC-tCS²⁵ and a device integrating the two monitors has been developed and characterized.^26,27 Despite of the successes of these studies, it must be noted that CW-fNIRS does not allow a quantification of absolute value of ${\mathrm{HbO}}_2$ and Hb concentration, lacks precision and accuracy,²⁸ and does not give a direct measure of CBF, the primary marker of the neurovascular coupling.

In this work, we propose to use time-resolved fNIRS (TR-fNIRS) and functional diffuse correlation spectroscopy (fDCS) as neuromonitors of the brain stimulation. Using pulsed light of few hundreds of picoseconds in width, TR-fNIRS measures absolute values of Hb and ${\mathrm{HbO}}_2$ concentration by separating the effects of absorption and scattering. The time information can also be used to explore different behaviors in different layers of the tissue.²⁹ On the other hand, fDCS³⁰ allows a direct measure of CBF in the microvasculature of the tissue.

Here, fDCS and TR-fNIRS are used together with EEG to monitor cerebral activity before, during, and after 10 min of DC-tCS over the frontal cortex. This area is the most convenient choice for the diffuse optical monitors because of the lack of hair. Since its stimulation has been proven effective for the improvement of various cognitive functions,^31–38 for example, working memory enhancement,^39–42 and for treatment of depression,^9,43,44 it is also a relevant area to stimulate and study.

2.

Methods

2.1.

Transcranial Direct Current Stimulation and Electroencephalography Device

DC-tCS was delivered using Neuroelectrics Starstim^® (Neuroelectrics, Barcelona, Spain), which is a research-class eight channel transcranial current stimulator that is also capable of measuring EEG. About 12-mm diameter silver chloride electrodes (Ag/AgCl) with conductive saline gel were used for EEG recording and for stimulation. Eight electrode (AF7, AF8, FT7, FT8, TP7, TP8, PO7, and PO8) locations were chosen, as depicted in Fig. 1(a), according to the 10–10 EEG system.⁴⁵ In particular, the active electrodes during stimulation were AF7 (left frontal lobe) and PO8 (right parieto-occipital lobe). From now on, we will refer to the former (AF7) as the stimulation electrode while to the latter (PO8) as the return electrode. All the eight electrodes functioned as EEG recording electrodes for protocol periods that did not involve stimulation. EEG spectrum was sampled at 500 Hz.

Fig. 1

(a) A schematic of the optical probes, stimulating and return electrodes and EEG recording electrodes superimposed on a standard 10–10 EEG system. (b) Image of the integrated probe that could accommodate optical fibers (both for fDCS and TR-fNIRS) and an electrode. (c) Schematic of the protocol that lasted a total of 50 min divided in five 10-min periods. Stimulation was delivered during the second period.

3.

Results

A total of 20 healthy subjects (nine female) participated in anodal and cathodal stimulation sessions. Twelve of them came back for a sham session. Technical problems resulted in one subject for TR-fNIRS measurement being discarded for the cathodal session leaving 19 subjects for ${\mathrm{HbO}}_2$ and Hb concentration measurements. In addition, one fDCS measurement was discarded during the sham session leaving 11 subjects for flow measurement during sham stimulation. Technical exclusion criteria for EEG, explained in Sec. 2.4, resulted in 20 subjects for anodal, 16 for cathodal, and 9 for the sham session. As explained above, outliers were detected and excluded from the statistical analysis. The number of subjects surviving the outlier detection for EEG and hemodynamics parameters is presented in Tables 1 and 2, respectively.

Table 1

The number of subjects for EEG that were used for the statistical analysis for each electrode and for each band, after the technical exclusion criteria and outlier detection.

Table 2

The number of subjects for hemodynamics parameters that were used for the statistical analysis after the technical exclusion criteria and outlier detection.

In Fig. 2, the time series of rCBF change, averaged over subjects, is shown together with the bootstrap sample. Only results from the left hemisphere (ipsilateral side, under the stimulation electrode) are shown from the three different stimulation sessions.

Fig. 2

rCBF change during the three types of stimulation sessions on the left hemisphere. (a) Sham session, (b) anodal session, and (c) cathodal session. The thick black line is the average, and the gray area represents the bootstrap sample over all the subjects. The black horizontal bar highlights the stimulation period.

When building the LME model from the null one, the stimulation type is found to improve it for blood flow in the left hemisphere, but not in the right one. Therefore, blood flow response depends on the stimulation in the ipsilateral hemisphere, but not in the contralateral one. Period is not an effect that improves the model for either of the hemispheres. Accordingly, we can conclude that blood flow stays constant at the end of the stimulation period (during) and in the post-periods. Posthoc analysis results are summarized in Fig. 6. Blood flow increases significantly from the baseline by 10% [confidence interval (CI): 7% to 14%] for anodal ($p<0.001$) and 11% (CI: 7% to 15%) for cathodal ($p<0.001$) stimulation while no increase is seen in the sham session ($p=0.07$). A comparison between the different sessions revealed a detectable difference between anodal and sham ($p=0.008$) and cathodal and sham ($p=0.003$) sessions while no difference is seen between anodal and cathodal sessions.

In Fig. 3, the results for the time evolution of ${\mathrm{HbO}}_2$ concentration from TR-fNIRS measurements using the three different analysis methods are shown for anodal and sham stimulations for the left (ipsilateral) hemisphere. DPF analysis gives an increase of ${\mathrm{HbO}}_2$ concentration starting during the stimulation and persisting after it. Notably, this change is seen both in the anodal and sham sessions. The results obtained with the analysis of variance and by fitting are comparable and show smaller changes than the DPF analysis both in anodal ($p<0.001$) and sham ($p<0.001$) sessions.

Fig. 3

Changes in oxyhemoglobin (${\mathrm{HbO}}_2$) concentration during (a) sham and (b) anodal stimulation, in the left hemisphere, according to the three types of analysis considered for TR-fNIRS: orange for DPF analysis, blue for the variance, and red for the fitting analysis. Continuous line is the average of all the subjects, the lighter colored region represents the bootstrap sample with the same color code. The black horizontal bar highlights the stimulation period.

For this reason, the fit analysis results are used for building the LME model for the changes in ${\mathrm{HbO}}_2$ and Hb concentrations. Furthermore, this is what is most commonly used for TR-fNIRS analysis.²⁹ Figures 4 and 5 show the average time evolution and the bootstrap sample from all the subjects for ${\mathrm{HbO}}_2$ and Hb concentrations, respectively, in the left hemisphere during the three stimulation sessions.

Fig. 4

Oxyhemoglobin (${\mathrm{HbO}}_2$) concentration during the three types of stimulation sessions on the left hemisphere from the fit analysis. (a) Sham session, (b) anodal session, and (c) cathodal session. The thick line is the average, and the lighter colored region represents the bootstrap sample over all the subjects. The black horizontal bar highlights the stimulation period.

Fig. 5

Deoxyhemoglobin (Hb) concentration during the three types of stimulation sessions on the left hemisphere from the fit analysis. (a) Sham session, (b) anodal session, and (c) cathodal session. The thick line is the average, and the lighter colored region represents the bootstrap sample over all the subjects. The black horizontal bar highlights the stimulation period.

As for the blood flow, among the effects considered in the LME model and listed in Sec. 2, only the stimulation improves the model. For both variables, the changes depend on the stimulation session but a recovery is not seen in the 30 min of post-periods, because the period is not a significant effect for the model.

As summarized in Fig. 6, ${\mathrm{HbO}}_2$ concentration increased by $0.5\mu \mathrm{M}$ (CI: 0.1 to $0.8\mu \mathrm{M}$) during the anodal session ($p=0.002$) and by $0.5\mu \mathrm{M}$ (CI: 0.2 to $0.8\mu \mathrm{M}$) during the cathodal ($p=0.001$) session. As expected, the changes in Hb concentration were opposite to ${\mathrm{HbO}}_2$ concentration. Hb concentration decreased by $-0.3\mu \mathrm{M}$ (CI: $-0.5$ to $-0.1\mu \mathrm{M}$) during the anodal session ($p<0.001$) and by $-0.3\mu \mathrm{M}$ (CI: $-0.5$ to $-0.2\mu \mathrm{M}$) during the cathodal session ($p<0.001$). No change was seen in the sham session ($p=0.1$ for ${\mathrm{HbO}}_2$ and $p=0.07$ for Hb concentration). A difference was detected by our model for ${\mathrm{HbO}}_2$ concentration between anodal and sham ($p=0.004$) sessions, as well as between cathodal and sham ($p=0.003$) sessions. On the other hand, a pair comparison for Hb concentration gave a significant difference only between cathodal and sham sessions ($p=0.03$).

Fig. 6

Changes of cerebral hemodynamics parameters during and after the stimulation sessions, as obtained by the LME model with the 95% confidence interval as the errorbars. $N$ refers to the number of subjects used for the statistical analysis. Red asterisk highlights statistically significant difference between stimulation sessions. Dashed vertical line represents the zero level.

A frequency range from 1 to 20 Hz was considered for the EEG analysis. The lower limit was chosen because artifacts derived from movement and subsequent changes in the skin-electrode impedance affect mostly the low EEG frequencies. Furthermore, a time–frequency analysis of each recording was performed in order to study the signal quality of the dataset. In many recordings, the EEG spectrum was affected by external interferences in frequencies above 22 Hz. It was then decided to choose an upper cutoff frequency of 20 Hz to avoid this noise. As a result, the beta band covers 13 to 20 Hz.

In addition, after the temporal and spectral inspection of the EEG spectrum that was acquired during the stimulation, we have concluded that the quality was not sufficient to perform a reliable EEG analysis. Therefore, only recordings before and after the stimulation were used.

Going through results obtained in EEG power, different changes from resting-state across stimulation modalities were observed only in the eyes-open condition. Figure 7 shows barplots for bands and electrodes where response was affected by the stimulation session according to LME model. Posthoc analysis resulted in a statistically significant power decrease in delta band in TP8 after every stimulation type. Statistically significant differences between anodal and cathodal sessions ($p<0.001$) and anodal and sham sessions ($p=0.02$) were observed. A statistically significant power decrease in the theta band after cathodal stimulation in fronto-right region (AF8, FT8) was observed. Statistically significant differences between cathodal and sham sessions ($p=0.002$ in both electrodes) and cathodal and anodal sessions ($p=0.002$ in both electrodes) were also observed. A more widespread effect in the beta band was observed. All electrodes in the tempo-parietal (TP7 and TP8) and parieto-occipital (PO7 and PO8) lobes experienced a statistically significant increase after sham, but, not after any active stimulation. In all these electrodes, the difference between anodal and sham sessions ($p=0.004$ for TP7, $p=0.01$ for TP8, $p<0.001$ for PO7, $p=0.002$ for PO8) and cathodal and sham sessions was significant ($p<0.001$ for all electrodes).

Fig. 7

Change of EEG power in electrode TP8 in the delta band, electrodes AF8 and FT8 in the theta band, and electrodes TP7, PO7, TP8, PO8 in the beta band caused by the three types of stimulation, as obtained by the LME model, with 95% confidence interval as the errorbars. $N$ refers to the number of subjects used for the statistical analysis. Red asterisk highlights the statistically significant difference between stimulation sessions. Dashed vertical line represents the zero level. Only bands and electrodes with statistical significant changes are shown.

4.

Discussion

In this work, we have introduced the use of fDCS, together with TR-fNIRS, as neuromonitors of cerebral hemodynamics during and after DC-tCS. The former has allowed the direct measurement of CBF primary biomarker of the neurovascular coupling. The measurement of cerebral hemodynamics was concurrent to EEG, which measures electrical activity. Therefore, information on both hemodynamics and neuronal activity from the same run of the protocol were available.

Before going through the results of the hemodynamics parameters, we note that the fDCS analysis requires the optical properties of the probed tissue as input parameters. In this experiment, the optical properties were measured by TR-fNIRS with fibers with a source–detector separation of 3.7 cm. On the other hand, source–detector separation of fDCS fibers was 2.5 cm, hence, probing a different volume than TR-fNIRS. Since our analysis assumes that the underlying tissue is homogeneous, we can tolerate this difference. It must also be highlighted that since we focus on the measurement of relative CBF, our results are relatively insensitive to small inaccuracies of the bulk optical properties.⁵⁶

Statistical analysis confirms that CBF increases in the left ipsilateral hemisphere (in the region under the stimulation electrode) during anodal and cathodal sessions but not during the sham session. CBF does not recover during the 30 min following the stimulation. This is consistent with the fact that if the stimulation is applied for a period longer than 9 min, the effect of the stimulation persists long after the stimulation ends.⁵⁷ Specifically, our statistical model results in a 10% increase during the anodal session and 11% during cathodal and gives no difference between the two stimulation conditions. This amount for CBF change is similar to what was assessed with arterial spin labeling, a magnetic resonance imaging technique,⁵⁸ but, is much larger than what measured by positron emission tomography (PET).⁵⁹ In Ref. 59, scans taken during the following hour after the stimulation were averaged, while we have measured only the following 30 min, which may account for some of the difference in our results. The optical probe placed in the right contralateral hemisphere did not detect any change in CBF in this region. The work done with PET⁵⁹ allowed for whole brain imaging and detected regional increases in different areas of the brain, even far away from the stimulation electrodes, after both anodal and cathodal sessions compared to sham. Although anodal stimulation proved to cause a more widespread increase of regional CBF than cathodal stimulation, the same effect was seen for both stimulation types in the region under the electrode, which is consistent with our results.

It is known that diffuse optical techniques measure a mixture of extra- and intracerebral perfusion.⁶⁰ This could be a confounding factor since DC-tCS causes a “skin redness” (erythema), which may lead to changes in the skin perfusion of different amounts between active and sham stimulations.¹⁰ However, the properties of photon diffusion into the tissue help us in explaining that response detected in this protocol is not mainly due to changes in skin perfusion due to the stimulation induced skin redness. The visitation probability of photons traveling into the tissue has a so-called “banana-shaped” pattern and touches the skin mainly in the region corresponding to the source and detector fibers.⁶¹ In our case, the fibers are separated by a distance of 2.5 cm while the electrode, placed in the middle, has a diameter of 1.2 cm. If we consider that various works^62,63 have shown that the erythema is concentrated in the area under the electrode, our optical monitors are not particularly sensitive to this redness. Furthermore, a recent study⁶⁴ measured the temperature under the electrode during the stimulation, detecting a nonsignificant change after 10 min of stimulation (our total time of stimulation). Only after 20 min of application, a significant increase of 1.36 °C was detected. However, we should account for the fact that this work used a larger electrode resulting in a lower current density than ours (0.06 versus $0.88{\mathrm{mA}/\mathrm{cm}}^2$). Interestingly, there is evidence suggesting that the current intensity and the electrode area are more important parameters leading to increased redness and discomfort due to the stimulation of more cutaneous receptors.⁶² In the above mentioned Ref. 64, temperature was measured using a current intensity of 2 mA, larger that what was used in our experiment. Therefore, we expect a minimal, nonsignificant temperature increase in our case. Another indication that the observed, long-term changes are not purely due to skin redness is that it is known that erythema disappears within 30 min after the stimulation.⁶⁵ In our case, the stimulation response did not decrease after 30 min.

Proceeding to TR-fNIRS measurements, results from the different analysis methods considered were compared. To start with, DPF analysis gives us the results that would have been obtained using a typical CW-fNIRS setup for fNIRS, the most commonly used fNIRS technology. DPF analysis shows an increase of about $2\mu \mathrm{M}$ in ${\mathrm{HbO}}_2$ concentration in the left hemisphere compatible with what was previously found in the literature.^20,22 This increase is seen in active and sham stimulations, where the brain is not actually stimulated which questions its validity. Focusing on the results given by the analysis of variance, the second moment of the DTOF, we retrieve smaller changes in both anodal and sham sessions with respect to DPF analysis. Notably, the variance of TR-fNIRS curves is more sensitive to the deeper layers with respect to the intensity, the parameter used in the DPF analysis.⁴⁸ This may suggest that results in DPF analysis are contaminated by superficial effects present both during anodal and sham stimulations. Moreover, these results are in accordance with a previous single subject experiment, where TR-fNIRS was used during anodal stimulation.²¹ In this work, TR-fNIRS measurements were analyzed with a gated analysis.⁶⁶ This method exploits the depth information encoded in time in the broadened pulse collected after propagation into the tissue and allows to decouple changes in the extracerebral and cerebral layers. Measurements performed on a single subject during anodal stimulation retrieved a larger ${\mathrm{HbO}}_2$ concentration change in the superficial than in the deeper layer, which would be heavily reflected in the DPF analysis. We would like to point out that the time gating analysis of TR-fNIRS was not implemented here because the lack of knowledge of the extracerebral layer thickness on an individual basis compromises the validity of this method.⁶⁷

The usage of TR-fNIRS has allowed us to improve the separation of the extracranial and intracranial signals for the analysis of the hemoglobin concentration changes. We do not have a comparable time-domain measurement for fDCS at the moment.^68,69 However, it has been proved that the relative brain-to-scalp sensitivity is about three times higher for fDCS than for CW-fNIRS.⁷⁰ Therefore, we believe that our results are reflective of the CBF in a more reliable manner than the CW-fNIRS results.

Since the fit analysis gives results comparable to the variance, and this is the most commonly used method when measuring oxygenation with TR-fNIRS, we have decided to consider the results given by the fit analysis.

In this protocol, ${\mathrm{HbO}}_2$ concentration increases and Hb concentration decreases in the left hemisphere during both anodal and cathodal stimulations, but not during sham. The right frontal lobe (contralateral to the stimulation) does not present changes in the oxygenation as is also true for CBF. Even if changes retrieved were very small, less than $1\mu \mathrm{M}$, ${\mathrm{HbO}}_2$ and Hb concentrations can mark the status of the ongoing stimulation. As a matter of fact, the changes are statistically significant during both anodal and cathodal sessions but not during the sham session. ${\mathrm{HbO}}_2$ concentration is a more robust marker, since response to anodal and to sham sessions is statistically significantly different, as well as the response to the cathodal and to sham sessions. On the other hand, for Hb concentration, the contrast-to-noise ratio is not enough to give a statistically significant difference between anodal and sham sessions, but a decrease is not seen in the sham session.

The directions of Hb and ${\mathrm{HbO}}_2$ concentration changes are consistent with what is expected in the case of a cerebral activation and what was previously found in the literature.^20,22 Specifically, due to the neurovascular coupling, when a cerebral region is activated, CBF is recruited to meet the additional demand of oxygen and it increases.¹⁷ Concurrent to this, an increase of ${\mathrm{HbO}}_2$ concentration and a decrease of Hb concentration are usually detected.⁷¹ This suggests that the brain is over supplied by oxygen by the increase of CBF compared to the oxygen consumption. The details of this process are, in fact, more complex.¹⁷ The time series of the hemodynamics parameters in Figs. 2, 4, and 5 suggest a steeper increase of CBF than the ${\mathrm{HbO}}_2$ and Hb concentrations. The above-mentioned mechanism of neurovascular coupling can explain this.⁷¹ Right after the stimulus, the supply of CBF does not overcome the need immediately. Only after this happens, a ${\mathrm{HbO}}_2$ concentration increase and Hb decrease is detected. This process was previously verified during in vivo protocols.⁷² Nonetheless, to make quantitative inferences on the time lag, we should have had a better time resolution than what we have here.

Both the anodal and cathodal stimulation sessions have created similar changes in the hemodynamics parameters. All the reviewed and cited published works^58,59 that measure cerebral hemodynamics during DC-tCS show the same pattern for both cathodal and anodal stimulations in the region under the electrode. Our results confirm these findings. In addition, PET studies concurrent to transcranial magnetic stimulation also demonstrate an increase in CBF in the region under the stimulation for both excitatory and inhibitory protocols.⁷³ These results suggest that CBF is not a marker of the polarity of the stimulation, at least in the directly stimulated area, and that regardless of whether the stimulation results in an activation of inhibitory or excitatory synapses, it leads to an increase in blood flow. It has been highlighted that this behavior of CBF reflects the local levels of synaptic activity in intracortical neurons and inputs to those areas rather than the activity of output neurons, which is opposite when the stimulation is excitatory or inhibitory.^59,74

The stimulation protocol used in this study influenced EEG rhythmic activity at different bands and electrodes. Comparing the results from previously published studies about DC-tCS on the frontal cortex is not a trivial task since EEG response depends heavily on the electrode montage.⁷⁵ All of the previous studies used F3 as the stimulation electrode while the return electrode, here referring to the electrode not placed in the frontal cortex with opposite polarity, was placed either on the supraorbital region or on extracerebral regions. By contrast, we have decided to use the AF7 electrode as stimulation electrode because the optical measurement series on the frontal brain region are more reliable due to the absence of hair. The return electrode was chosen to be spatially and functionally distant from the stimulation electrode in order to minimize the interactions and to stimulate the frontal region unilaterally.

Moreover, few studies measured the EEG spectrum at rest while most of them assessed the response to a task and how this can be influenced by the stimulation. It must be noted that a recent systematic review highlighted that results in the EEG spectrum of the same protocol were not consistent within different works.⁷⁶ They concluded that there is no evidence that the EEG spectrum is affected by the stimulation. It may also be hypothesized that, since it is not trivial to maintain subjects in a well-controlled condition, the variability of the effects is increased.

We note that the protocol is 50 min long, which may have led to a sense of drowsiness and/or fatigue in the subjects, which could alter the cerebral electrophysiology and metabolism. The staring and attention tasks were implemented to minimize this effect. Furthermore, the transition from eyes open and closed and vice versa also helped with this issue. In spite of this, we cannot exclude that some of the subjects may have been affected and we do not have any control over this.

In order to define EEG spectrum results, effects in different frequency bands have been studied since different bands are traditionally related to different brain functions. It is, however, controversial whether the activity at a certain frequency band is related to a single brain function or that complex stimuli are reflected on a single oscillatory activity.⁷⁷ This may explain the multiplicity of bands affected by our protocol.

The protocol applied in this study influenced brain activity in the delta, theta, and beta bands. No effect was detected in the alpha band, which is consistent with previous works.⁷⁶

We will focus, from now on, on the bands, where the stimulation has proven to be effective. The changes in the delta band were observed in the right tempo-parietal region after all stimulation sessions but were significantly smaller in the anodal session compared to the other two. On the other hand, changes in the theta band occurred in the right fronto-temporal area exclusively after cathodal stimulation. Interestingly, delta and theta bands are associated with functions linked to the frontal cortex, which was stimulated in our protocol. It is an indication of activity in cognitive and memory performances,⁷⁸ attention,⁷⁹ and working memory.⁸⁰ A response in these bands is consistent with previous work. Specifically, a lower power in delta and beta bands after the stimulation compared to sham has been previously observed,^35,81 where pre- and post-stimulation spectral power was not calculated but the comparison between stimulation sessions was done in absolute values. This effect has been found only in the frontal cortex, in contrast to our findings, but both stimulation and return electrodes were placed on the anterior part of the head, while we have an electric field crossing all the brain. A significant and selective decrease of the power of theta band in the right inferior frontal gyrus area has been reported after DC-tCS stimulation along with an improvement in behavioral inhibition.⁸²

Differences in EEG power across different stimulation sessions were also seen in the beta band. Power increased after the sham session in temporal and parietal regions, but not after active stimulation sessions. An unchanged or a decreased level of beta activity is usually observed in tasks where behavior is driven by bottom-up signals, while enhancement occurs if the system has to actively maintain endogenous attention.⁸³ It may be hypothesized that the detected increase in beta activity is linked to an active effort to maintain the cognitive concentration state required for the sustained attentional task along the experimental protocol. This activity was present during sham but not during active stimulation sessions. Several references highlight an influence of DC-tCS when applied on the frontal cortex, on sustained attention.⁸⁴ In our protocols, the same effects for anodal and cathodal sessions were found in the EEG beta band in the posterior part of the brain, without a differentiation due to the polarity. Intuitively, this is not expected because opposite polarity stimulation should produce an opposite electric field in the brain, i.e., either excitatory or inhibitory. Nonetheless, it must be mentioned that the predominance of excitatory or inhibitory electric field in an area of the brain does not exclude that some “pockets” of the opposite direction are present.⁸⁵ We can speculate that only excitatory areas influenced the power in the beta band in the posterior area of the brain in our protocol. Unfortunately, this is prohibitively hard to confirm since not only the electrode montage but also the concurrent tasks alter the response to a certain polarity of stimulation,^86,87 i.e., it would require a special protocol to address this. For our protocol, the attention task was not the primary goal.

Last, it was verified whether the change in CBF in the left frontal cortex due to stimulation has any correlation with the power change respect to baseline in any of the EEG bands and electrodes. This was done even for the electrodes far away from the optical probes since the source of the EEG signal can lay far from the scalp point, where it is registered and its exact location cannot be well determined due to the inverse problem.⁸⁸ The Pearson correlation coefficient was calculated and a correlation was defined significant if, by bootstrapping the sample (i.e., repeating the analysis removing one subject at time), the correlation stays significant. After this analysis, we have found a negative correlation in the delta band for electrode FT8 with a Pearson correlation coefficient of $R=-0.6$. This is consistent with a predicted and verified decrease in lower frequency band during activation.^89,90 The study of the correlation between the hemodynamics and neuronal activity is a very appealing topic due to the important information on the neurovascular coupling it could provide. It is also a very complicated process since several factors can contribute to the failure of the correlation between EEG activity and hemodynamics.⁹¹ A detailed characterization of this is beyond the scope for our paper, which aimed at the introduction of hybrid diffuse optical devices for monitoring the response to the stimulation, and requires a dedicated protocol with more regions probed for hemodynamics.

In summary, we have introduced fDCS as a neuromonitoring technique for following the response to DC-tCS. It was integrated with TR-fNIRS to measure the concentrations of ${\mathrm{HbO}}_2$ and Hb in a more accurate way with respect to CW-fNIRS. We have proved that fDCS and TR-fNIRS can be applied concurrently to EEG measurements, which assess cerebral activity. We were able to have information on both cerebral hemodynamics and neuronal activity in the same run of stimulation.

5.

Conclusion

We have shown that fDCS and TR-fNIRS are suitable tools to monitor cerebral hemodynamics during and after DC-tCS and they can be integrated with EEG, which monitors the neuronal activity. We have found CBF to be a good indicator of the ongoing stimulation since it showed an increase during and after anodal and cathodal stimulation in the region under the stimulation electrode. Results obtained with TR-fNIRS showed small changes of ${\mathrm{HbO}}_2$ and Hb concentrations but the effect of the stimulation was still detectable.

By concurrent EEG recordings, we could follow the modulation of the underlying neural oscillations. DC-tCS over the frontal cortex induced statistically significant power changes across different stimulation sessions in delta, theta, and beta EEG rhythms.