Electromagnetic-based methods of stimulating brain activity require invasive procedures or have other limitations. Deep-brain stimulation requires surgically implanted electrodes. Transcranial magnetic stimulation does not require surgery, but suffers from low spatial resolution. Optogenetic-based approaches have unrivaled spatial precision, but require genetic manipulation. In search of a potential solution to these limitations, we began investigating the influence of transcranial pulsed ultrasound on neuronal activity in the intact mouse brain. In motor cortex, ultrasound-stimulated neuronal activity was sufficient to evoke motor behaviors. Deeper in subcortical circuits, we used targeted transcranial ultrasound to stimulate neuronal activity and synchronous oscillations in the intact hippocampus. We found that ultrasound triggers TTX-sensitive neuronal activity in the absence of a rise in brain temperature (<0.01°C). Here, we also report that transcranial pulsed ultrasound for intact brain circuit stimulation has a lateral spatial resolution of approximately 2 mm and does not require exogenous factors or surgical invasion.

Ultrasound can be defined as low or high intensity (). High-intensity US (>1 W/cm) influences neuronal excitability by producing thermal effects (). In addition to the initial studies cited above, high-intensity US has been reported to modulate neuronal activity in peripheral nerves (), craniotomized cat and craniotomized rabbit cortex (), peripheral somatosensory receptors in humans (), cat spinal cord (), and rodent hippocampal slices (). While these prior studies support the general potential of US for neurostimulation, high-intensity US can readily produce mechanical and/or thermal tissue damage (), precluding it from use in noninvasive brain-circuit stimulation. At acoustic intensities <500 mW/cm, pulsed US can produce mechanical bioeffects without producing thermal effects or tissue damage (). In hippocampal slices, we previously reported low-intensity US (<300 mW/cm), low-frequency US (<0.65 MHz) is capable of stimulating action potentials and synaptic transmission (). Since low-frequency US can be reliably transmitted through skull bone (), the motivation for the present study was to investigate the influence of low-frequency, low-intensity transcranial pulsed US on intact brain circuits in pursuit of a novel brain-stimulation method. We report that transcranial US is capable of safely and reliably stimulating in vivo brain circuits, such as the motor cortex and intact hippocampus of mice.

The significance of membrane changes in the safe and effective use of therapeutic and diagnostic ultrasound.

Edmund Newton Harvey first published a set of ground-breaking observations that clearly described that US can stimulate nerve and muscle fibers in neuromuscular preparations (). Since then, US has been shown to stimulate and inhibit neuronal activity under various conditions. For example, US has been reported to reversibly suppress sensory-evoked potentials in the cat primary visual cortex following treatment of the lateral geniculate nucleus with US transmitted through a cranial window (). Conversely, US has been shown to stimulate auditory nerve responses in the craniotomized cat brain (). In cat saphenous nerve preparations, US was shown to differentially modulate the activity of Aδ- and C-fibers, depending on the fiber diameter, US intensity, and US exposure time ().

Ultrasound is a mechanical pressure wave (sound wave) having a frequency above the range of human hearing (>20 kHz). Due to its physical properties, specifically its ability to be transmitted long distances through solid structures, including bone and soft tissues, US is used in a wide range of medical and industrial applications. Diagnostic imaging US has a frequency range from 1 to 15 MHz, while therapeutic US tends to employ a frequency of about 1 MHz (). Ultrasound can be transmitted into tissues in either pulsed or continuous waveforms and can influence physiological activity through thermal and/or nonthermal (mechanical) mechanisms (). The potential of using US for brain stimulation has been largely overlooked in comparison to chemical, electrical, magnetic, or photonic methods. Surprisingly, this is in lieu of the fact that US was shown capable of exciting nerve and muscle more than eight decades ago ().

The significance of membrane changes in the safe and effective use of therapeutic and diagnostic ultrasound.

All currently implemented approaches to the stimulation of brain circuits suffer from a limitation or weakness. Pharmacological and chemical methods lack brain target specificity and have numerous metabolic requirements. Electrical methods, such as deep-brain stimulation, offer a higher targeting specificity but require surgery and brain impalement with electrodes (). Optogenetic-based methods using light-activated ion channels or transporters offer unrivaled spatial resolution but require genetic alteration (). Transcranial magnetic stimulation (TMS) and transcranial direct current stimulation do not require invasive procedures but suffer from poor spatial resolution of ≈1 cm (). Considering the above limitations, a remaining challenge for neuroscience is to develop improved stimulation methods for use in intact brains. To address this need, we began studying the influence of pulsed ultrasound (US) on neuronal activity in mice.

We naturally questioned whether these effects were accompanied by the regulation of activity-mediated cellular molecular signaling cascades in the hippocampus. Brain-derived neurotrophic factor (BDNF) is one of the most potent neuromodulators of hippocampal plasticity, and its expression/secretion is known to be regulated by neuronal activity (). We thus examined BDNF protein expression levels in the hippocampus following transcranial stimulation with pulsed US. Unilateral hippocampi of mice (n = 7) were targeted and stimulated with pulsed US (0.35 MHz, 50 cycles per pulse, 1.5 kHz PRF, 500 pulses) having an I= 36.20 mW/cmevery 2 s for 30 min. Following a 45 min recovery period, mice were sacrificed and their brains removed, sectioned, and immunolabeled with antibodies against BDNF. We observed that pulsed US induced a significant increase in the density of BDNFpuncta in CA1 s.p. (contralateral control = 149.64 ± 11.49 BDNFpuncta/7.5 × 10mmfrom 0.61 mmCA1 region/mouse versus US stim = 221.50 ± 8.75 BDNFpuncta/7.5 × 10mmfrom 0.61 mmCA1 region/mouse; t test, p < 0.001; Figure 7 D). Similar significant increases were observed in the CA3 s.p. region (contralateral control = 206.20 ± 19.68 BDNFpuncta/7.5 × 10mmfrom 0.61 mmCA3 region/mouse versus US stim = 324.82 ± 27.94 BDNFpuncta/7.5 × 10mmfrom 0.61 mmCA3 region/mouse; t test, p < 0.005; Figure 7 D). These data demonstrate that pulsed US can be used to remotely stimulate neuronal activity in the intact mouse hippocampus. Posing captivating potential for broad applications in neuroscience, the increased synchronous activity and elevated BDNF expression patterns produced by pulsed US lend support to our hypothesis that transcranial US can be used to promote endogenous brain plasticity.

Pulsed US produced a significant (p < 0.01) increase in spike frequency lasting 1.73 ± 0.12 s ( Figure 7 B). Natural activity patterns in the CA1 region of hippocampus exhibit gamma (40–100 Hz), sharp-wave (SPW) “ripple” (160–200 Hz), and other frequency-band oscillations reflecting specific behavioral states of an animal (). Sharp-wave ripples (≈20 ms oscillations at ≈200 Hz) in CA1 result from the synchronized bursting of small populations of CA1 pyramidal neurons () and have recently been shown to underlie memory storage in behaving rodents (). On the other hand, the consequences of gamma oscillations in the CA1 region of the hippocampus are not as well understood but are believed to stem from the intrinsic oscillatory properties of inhibitory interneurons (). By decomposing the frequency components of wideband (1–10,000 Hz) activity patterns evoked by pulsed US, we found that all after-discharges contained both gamma oscillations and SWP ripple oscillations lasting <3 s ( Figure 7 C and S5 ). These data demonstrate that pulsed US can stimulate intact mouse hippocampus while evoking synchronous activity patterns and network oscillations; hallmark features of intrinsic hippocampal circuitry.

We used an angled line of US transmission through the brain by positioning acoustic collimators 50° from a vertical axis along the sagittal plane. The output aperture of collimators (d = 2 mm) were unilaterally centered over −4.5 mm of Bregma and 1.5 mm lateral of the midline ( Figure 7 A ). We used a 30° approach angle to drive tungsten microelectrodes to the CA1 s.p. region of hippocampus through cranial windows (d = 1.5 mm) centered approximately −1.0 mm of Bregma ( Figure 7 A). Pulsed US (0.25 MHz, 40 cycles per pulse, 2.0 kHz PRF, 650 pulses) having an I= 84.32 mW/cmreliably triggered an initial LFP with a mean amplitude of −168.94 ± 0.04 μV (50 trials each) and a mean response latency of 123.24 ± 4.44 ms following stimulus onset ( Figure 7 B and S5 ). This initial LFP was followed by a period of after-discharge activity lasting <3 s ( Figure 7 B and S5 ). These short-lived after-discharges did not appear to reflect abnormal circuit activity as observed during epileptogenesis (). In fact, hippocampal after-discharges lasting more than 10 s are indicative of seizure activity ().

(D) Confocal images illustrating BDNF (green) expression in the CA1 s.p. (top) and CA3 s.p. (bottom) regions of hippocampus from contralateral control (left) and stimulated hemispheres (right). Histograms (far right) illustrate the significant increase in the density of BDNF + puncta triggered by transcranial US stimulation for the CA1 s.p. (top) and CA3 s.p. (bottom) regions of hippocampus. Data shown are mean ± SEM.

(C) An individual recording trace of CA1 s.p. extracellular activity in response to a pulsed US waveform is shown in its wideband (top), gamma (middle), and SWP (bottom) frequency bands. An expanded 250 ms region of the SWP trace (red) illustrates SWP “ripples” (also see Figure S5 ).

(B) Raw (black) and average (cyan) hippocampal CA1 LFP recorded in response to 50 consecutive US stimulation trials (left). A psuedocolored spike-density plot illustrates the increase in CA1 s.p. spiking as a function of time in response to 50 consecutive pulsed US stimuli delivered at 0.1 Hz (right).

(A) Shown is an illustration of the geometrical configuration used for targeting the dorsolateral hippocampus with transcranial pulsed US while recording evoked electrophysiological responses in the dorsal hippocampus (left). A lesion illustrates the site of an electrophysiological recording location in the hippocampal CA1 s.p. region (right).

We finally aimed to determine if trancranial pulsed US can be used to stimulate subcortical brain circuits in intact mice. To address this issue, we focused our attention on the intact mouse hippocampus, since pulsed US waveforms have been shown to elicit action potentials and synaptic transmission in hippocampal slices (). We performed extracellular recordings of US-evoked activity in the CA1 stratum pyramidale (s.p.) cell body layer of dorsal hippocampus (n = 7 mice). Prompted by our observations regarding the potential disruption of US fields by dense white matter tracts, we implemented a targeting approach bypassing the dense white matter of the corpus callosum when transmitting pulsed US to the hippocampus.

Through our development of the US brain-stimulation method described above, we have stimulated the intact brains of more than 190 mice through >92,000 US stimulus trails. We allowed >50% of the mice to recover from anesthesia following stimulation procedures and never observed any neurological abnormalities such as paralysis, ataxia, or tremor in these mice. Even mice undergoing multiple repeated-stimulation protocols spanning a 2 week time period ( Figure S4 A) exhibited no visible behavioral impairments or signs of diminishing responsiveness to transcranial pulsed US. In our studies, fewer than 6% of the animals died during or immediately following a US stimulation experiment. This mortality rate was likely due to respiratory or cardiac complications associated with maintaining mice under ketamine/xylazine anesthesia for extended periods of time (>2 hr). Based on the collective observations described above, we conclude that low-intensity transcranial pulsed US provides a safe and noninvasive method of stimulating intact brain circuit activity in mice. Whether similar safety margins hold true for other animal species must be directly evaluated and remains undetermined.

To determine if transcranial US stimulation of motor cortex produced any gross impairments in motor behavior. The day before stimulation with pulsed US waveforms (I= 142.2 mW/cm; every 10 s for 30 min), 24 hr poststimulation, and again 7 days poststimulation, we performed a series of experiments designed to assay motor function. Compared to sham-treated controls (n = 9 mice), a repeated-measures ANOVA revealed no significant effect of US stimulation (n = 9 mice) on a rotorod running task (F= 0.211, p > 0.1; Figure S4 C). We also measured motor function and grip strength by subjecting mice to a wire-hanging task. Again, repeated-measures ANOVA revealed no significant group effect on hang time (F= 0.05; p > 0.1; Figure S4 C). During daily behavioral monitoring, we observed no differences in feeding behavior, grooming behavior, or startle reflexes between US-stimulated mice and sham controls.

To determine the effects of pulsed US on brain ultrastructure, we used quantitative transmission electron microscopy to examine stimulated and control brains. We compared excitatory synapses in the motor cortex from control unstimulated mice (n = 5 mice) with synapses in the stimulated regions of motor cortex from mice (n = 6) that underwent a US stimulus trial as described above (I= 142.2 mW/cm) every 10 s for 30 min ( Figure 6 E). An independent samples t test revealed no significant difference in the density of synapses between groups (control = 16.59 ± 0.81 synapses/100 μmfrom 2.3 mmversus US stim = 22.99 ± 4.07 synapses/100 μmfrom 4.2 mm; p > 0.10; Figure 6 F). There were also no significant differences in the postsynaptic density (PSD) length (control = 0.225 ± 0.009 μm from 99 synapses versus US stim = 0.234 ± 0.009 μm from 130 synapses; p > 0.10), the area of presynaptic terminals (control = 0.279 ± 0.02 μmversus US stim = 0.297 ± 0.02 μm; p > 0.10), the density of vesicles in presynaptic boutons (control = 206.89 ± 9.52 vesicles/μmversus US stim = 209.85 ± 8.14 vesicles/μm; p > 0.10), or the number of docked vesicles (DV) occupying active zones (control = 21.71 ± 0.91 DV/μm versus US stim = 20.26 ± 0.61 DV/μm; p > 0.10) between treatment groups ( Figure 6 F). There were no qualitative differences in the ultrastructure of cortical neuropil between treatment groups ( Figure S4 B).

We next probed the cellular-level consequences of pulsed US on brain tissues using antibodies against cleaved caspase-3 to monitor cell death ( Figure 6 C). Using the same US waveform described above (I= 142.2 mW/cm), we unilaterally stimulated the motor cortex of mice (n = 8) every 10 s for 30 min. Following a 24 hr recovery period to allow for peak caspase-3 activation, mice were sacrificed and their brains examined using confocal microscopy. In comparing stimulated cortex regions with their contralateral controls (2.81 mmtotal area/hemisphere/mouse), we found that pulsed US did not induce a change in the density of apoptotic glial cells (control = 0.40 ± 0.04 caspase-3cells/0.56 mmversus US Stim = 0.43 ± 0.06 caspase-3cells/0.56 mm; p > 0.30) or apoptotic neurons (control = 0.08 ± 0.03 caspase-3cells/0.56 mmversus US stim = 0.07 ± 0.03 caspase-3cells/0.56 mm; p > 0.50; Figure 6 D). To further confirm this lack of an effect on cell death, we repeated the above experiment in mice (n = 4) using a higher-intensity US waveform (I= 300 mW/cm), which is 137 mW/cmhigher intensity than we used to evoke brain activity with any waveform in this study. We again observed no significant effects (2.81 mmtotal area/hemisphere/mouse) of pulsed US on the density of apoptotic glial cells (control = 0.44 ± 0.16 caspase-3cells/0.56 mmversus US stim = 0.38 ± 0.13 caspase-3cells/0.56 mm; p > 0.30) or apoptotic neurons (control = 0.06 ± 0.05 caspase-3cells/0.56 mmversus US stim = 0.07 ± 0.05 caspase-3cells/0.56 mm; p > 0.50; Figure 6 D).

To assess the safety of transcranial US brain stimulation in mice, we first examined how pulsed US influenced blood-brain barrier (BBB) integrity. Prior to stimulation, mice received an intravenous administration of fluorescein isothiocyanate-dextran (10 kDa), which does not cross the BBB under normal conditions (). The motor cortex of mice (n = 5) was then unilaterally stimulated every 10 s for 30 min with pulsed US (0.50 MHz, 225 cycles per pulse, 1.5 kHz PRF, 100 pulses) having an I= 142.20 mW/cmusing a collimator (d = 4.7 mm). We observed no evidence that US produced damage to the BBB, as indicated by a complete lack of fluorescein leakage (contralateral control = 179.6 mm vasculature length examined versus US Stim = 183.4 mm vasculature length examined; Figure 6 A ). In separate positive control experiments, we coadministered intravenous fluorescein-dextran with an US contrast agent (Optison) shown to mediate in vivo BBB disruption in response to US (). Results from these positive control experiments (n = 3 mice) confirmed our ability to detect BBB damage had it occurred in response to pulsed US alone ( Figure 6 B).

(F) Histograms are shown for mean synaptic density (top left), mean axonal bouton synaptic vesicle density (top right), mean PSD length (bottom left), and mean number of DV occupying active zones (bottom right).

(D) Histograms illustrate the mean density of cleaved caspase-3 + glial cells (“G”) and neurons (“N”) observed in the motor cortex of contralateral control and US-stimulated hemispheres for two different stimulus intensity waveforms. Data shown are mean ± SEM.

(C) Confocal images of NeuN + (green) and cleaved caspase-3 + (magenta) cells obtained from a US-stimulated region show positive glial cells (top) and a neuron (bottom) at low- (left) and high magnification (right).

(B) A similar confocal image is shown, but was obtained from a positive control treatment group where US-stimulation was performed in the presence of Optison, an ultrasound-microbubble contrast agent known to elicit cavitationally mediated vasculature damage.

(A) Confocal images of TO-PRO-3-labeled cells (red) and fluorescein-dextran-filled cerebrovasculature (green) obtained from the motor cortex of a contralateral control hemisphere (left) and from the stimulated region of the US-treated hemisphere (right).

By examining the effects of pulsed US along the dorsal-ventral axis within the stimulation zone (0.5–2.5 mm medial to lateral; −1.2 to −3.2 mm of Bregma), we found the density of c-foscells was significantly higher (p < 0.05) compared to contralateral controls in the superficial 1.0 mm of tissue ( Figure S3 C). While there were trends of higher c-foscell densities in some deeper nuclei of stimulated hemispheres, we only observed one significant difference in a deep-brain region ( Figure S3 C). The elevated c-fos here may have been produced by standing waves or reflections, since higher c-foscell densities were generally observed near the skull base. Otherwise, we would have expected to observe elevated c-foslevels uniformly along the dorsal-ventral axis of stimulated regions due to the transmission/absorption properties of US in brain tissue. For >1.5 mm of the 2.0 mm diameter cortical area we targeted with US in these mapping studies, regions deeper than ≈1 mm were ventral to dense white matter tracts (corpus callosum) in the brain. Interestingly, unmyelinated C-fibers have been shown to be more sensitive to US than myelinated Aδ fibers (). Effectively blocking US-evoked activity in subcortical regions, we suspect low-intensity US fields may have been absorbed/scattered by dense white matter tracts in these mapping studies as a function of the US transmission path implemented. Despite these observations, we show below that it is indeed possible to stimulate subcortical brain regions with transcranial US by employing different targeting approaches (see Remote Stimulation of the Intact Hippocampus Using Transcranial Pulsed US).

We prepared coronal sections from brain regions spanning +0.25 mm to −4.20 mm of Bregma ( Figure 5 A). Individual sections spaced every 125 μm were then immunolabeled using antibodies against c-fos and imaged using transmitted light microscopy. We quantified c-foscell densities in 250 × 250 μm squares for entire coronal sections, corrected for tissue shrinkage, and developed brain activity maps by plotting c-foscell densities in 250 × 250 μm pixels onto their corresponding anatomical locations using mouse brain atlas plates (). Representative raw data and functional activity maps coding c-foscell density using a psuedocolor lookup table for visualization purposes are shown in Figures 5 B–5D. We estimated the lateral resolution of pulsed US along the rostral-caudal brain axis by analyzing regions of dorsal cortex (0.25–1.0 mm deep; 0.75–1.50 mm lateral of the midline) for each coronal section ( Figures 5 A–5D). An ANOVA comparing the mean c-foscell densities for each 250 × 250 μm square region collapsed across animals revealed that pulsed US produced a significant increase in the density of c-foscells (ANOVA, F= 73.39, p < 0.001; contralateral control hemisphere mean c-foscell density = 16.29 ± 0.20 cells/6.25 × 10mmcompared to US stim = 19.82 ± 0.36 cells/6.25 × 10mm). Subsequent pairwise comparisons of stimulated versus contralateral control cortex revealed that US stimulation produced a significant increase in c-foscell densities for a 1.5 mm region along the rostral-caudal axis (−1.38 mm to −2.88 mm of Bregma) under the 2.0 mm diameter stimulation zone ( Figure 5 E). Similar analyses along the medial-lateral axis of dorsal cortex revealed a significant increase (p < 0.05) in c-foscell densities for a 2.0 mm wide region of brain tissue under the stimulation zone ( Figure S3 B). We observed a smearing of elevated c-foscell densities lateral to the stimulation zone, which could be attributed to nonlinearities in our acoustic collimators ( Figure S2 C), the corticocortical lateral spread of activity, and/or slight lateral variations in the positioning of our collimators.

To characterize the spatial distribution of US-evoked activity, we constructed functional activity maps using antibodies against c-fos (n = 4 mice). To facilitate data interpretation, we chose to stimulate intact brain tissue having a relatively planar surface and prominent subcortical structures. We centered the output of acoustic collimators (d = 2 mm; Figure S2 C) over the skull covering the right hemisphere from −1.2 mm to −3.2 mm of Bregma and 0.5 mm to 2.5 mm lateral of the midline using stereotactic coordinates ( Figure 5 A ;). We used our smallest-diameter collimator to characterize the minimal resolution of our brain-stimulation method since it is expected that larger collimators will produce larger areas of brain activation. Pulsed US (0.35 MHz, 50 c/p, 1.5 kHz PRF, 500 pulses) having an I= 36.20 mW/cmwas transmitted along a vertical axis parallel to the sagittal plane through underlying brain regions once every 2 s for 30 min. Following a 45 min recovery period, mice were sacrificed and their brains were harvested for histology.

(E) The line plots illustrate the mean c-fos + cell densities observed along the rostral-caudal axis of reconstructed brain volumes for stimulated (black) and contralateral control hemispheres (gray). Regions of cortex within the stimulation zone are indicated in red. Data shown are mean ± SEM.

(C) A psuedocolored map of c-fos + cell densities in 250 × 250 μm regions is shown for a reconstructed coronal section obtained from within the stimulus zone. Small regions inside (i) and outside (ii and iii) the US brain transmission path are highlighted and contain c-fos density data obtained from the corresponding images shown in (B).

(B) Light micrographs showing c-fos activity in a coronal brain section at different locations inside (i) and outside (ii and iii) the US transmission path.

(A) Diagrams showing the anatomical locations where transcranial pulsed US was delivered through an acoustic collimator (green; d = 2 mm; Figure S2 C) and the brain volume subsequently reconstructed (blue) to develop functional activity maps using antibodies against c-fos ( Figure S3 ).

We next examined how acoustic frequencies and intensities across the ranges studied here influenced US-evoked EMG responses from the triceps brachii of mice (n = 20). We stimulated motor cortex using 20 distinct pulsed US waveforms composed with different US frequencies (0.25, 0.35, 0.425, and 0.5 MHz) and having varied intensities ( Table S1 ). We randomized the sequence of which different waveforms were used in individual stimulus trials to avoid order effects. Relative comparisons of EMG amplitudes across animals can be influenced by many factors, including electrode placement, number of fibers recorded from, variation in noise levels, and differential fiber recruitment, which can be handled using normalization techniques to reduce intersubject variability (). To examine US-evoked EMG responses having the same dynamic range across animals, we normalized the peak amplitude of individual EMG responses to the maximum-peak amplitude EMG obtained for an animal and forced its minimum-peak amplitude EMG response through zero. A two-way ANOVA revealed a significant main effect of US frequency on EMG amplitude, where lower frequencies produced more robust EMG responses (F= 3.95, p < 0.01; Figure 4 A ). The two-way ANOVA also revealed a significant main effect of intensity (I) on EMG amplitudes (F= 9.78, p < 0.001; Figure 4 B), indicating that lower intensities triggered more robust EMG responses. The two-way ANOVA also revealed a significant frequency × intensity interaction (F= 7.25, p < 0.01; Figure 4 C), indicating differential effects of US waveforms on neuronal activity as a function of frequency and intensity. Across the stimulus waveforms studied, we found that the EMG response latencies were not affected by either frequency or intensity (data not shown).

(C) The interaction between US intensity (I SPTA ) and US frequency is plotted as a function of maximum-peak normalized EMG amplitudes (pseudocolor LUT).

(A) Maximum-peak normalized (Norm) US-evoked EMG amplitude histograms are plotted for the four US frequencies used in the construction of stimulus waveforms. Data shown are mean ± SEM.

We observed that application of TTX to motor cortex blocked EMG activity, which indicates that pulsed US triggers cortical action potentials to drive peripheral muscle contractions (n = 4 mice; Figure 3 D). The intensities of US stimuli we studied were <500 mW/cm, where mechanical bioeffects have been well documented in the absence of thermal effects (). To confirm these observations in brain tissue, we monitored the temperature of motor cortex in response to US waveforms having different pulse duration (PD) times. Equations for estimating thermal absorption of US in biological tissues indicate that PD times are a critical factor for heat generation () and predict that 0.5 MHz US pulses exerting a pof 0.097 MPa for a PD of 0.57 ms should produce a temperature increase of 2.8 × 10°C in brain (see Experimental Procedures ). All US stimulus waveforms used in this study had pvalues <0.097 MPa and PD times ≤0.57 ms. None of the US waveforms used to stimulate cortex elicited a significant change in cortical temperature within our 0.01°C resolution limits ( Figure 3 E). We found that US pulses with pvalues of 0.1 MPa and PD times >50 ms were required to produce a nominal temperature change (ΔT) of 0.02°C ( Figure 3 E).

The significance of membrane changes in the safe and effective use of therapeutic and diagnostic ultrasound.

By examining EMG failure rates in eight mice, we next studied how the success of achieving motor activation was affected when stimulus trials were repeated in more rapid succession. The mean EMG failure probability significantly increased (p < 0.001) as the rate of US stimulus delivery increased from 0.25 to 5 Hz ( Figure 3 C and Movie S3 ). These data suggest that brain stimulation with US may not be useful at stimulation frequencies above 5 Hz. To confirm these observations and further explore this potential limitation, future investigations of an expanded US stimulus waveform space are required because it is not known how other US waveform profiles will influence the generation of sustained activity patterns.

The baseline failure rate in obtaining US-evoked motor responses was <5% when multiple stimulus trials were repeated once every 4–10 s for time periods up to 50 min ( Figure 3 B). As observed for response latencies in acute experiments, the peak amplitudes of EMG responses evoked by transcranial pulsed US were stable across trial number ( Figure 3 B). In more chronic situations, we performed repeated US stimulation experiments within individual subjects (n = 5 mice) on days 0, 7, and 14 using a trial repetition frequency of 0.1 Hz for 12–15 min each day. In these experiments, there were no differences in the peak amplitudes of the US-evoked EMG responses across days (day 0 mean peak EMG amplitude = 40.26 ± 0.99 μV, day 7 = 43.06 ± 1.52 μV, day 14 = 42.50 ± 1.42 μV; ANOVA F= 1.47, p = 0.23; Figure S4 A). These data demonstrate the ability of transcranial US to successfully stimulate brain circuit activity across multiple time periods spanning minutes ( Figure 3 B) to weeks ( Figure S4 A).

When bilaterally targeted to motor cortex, pulsed US (0.50 MHz, 100 cycles per pulse, 1.5 kHz PRF, 80 pulses) having an I= 64.53 mW/cmtriggered tail twitches and EMG activity in the lumbosacrocaudalis dorsalis lateralis muscle with a mean response latency of 22.65 ± 1.70 ms (n = 26 mice). When unilaterally transmitted to targeted regions of motor cortex using a collimator (d = 3 mm), pulsed US (0.35 MHz, 80 c/p, 2.5 kHz PRF, 150 pulses) having an I= 42.90 mW/cmtriggered an EMG response in the contralateral triceps brachii muscle with a mean response with latency of 20.88 ± 1.46 ms (n = 17 mice). With nearly identical response latencies (21.29 ± 1.58 ms), activation of the ipsilateral triceps brachii was also observed in ∼70% of these unilateral stimulation cases ( Movie S2 ). Although consistent from trial to trial ( Figure 3 B), the EMG response latencies produced by US brain stimulation were ≈10 ms slower than those obtained using optogenetic methods and intracranial electrodes to stimulate motor cortex (). Several reports show that TMS also produces response latencies slower than those obtained with intracranial electrodes (). Discrepancies among the response latencies observed between electrical and US methods of brain stimulation are possibly due to differences in the time-varying energy profiles that these methods impart on brain circuits. The underlying core mechanisms of action responsible for mediating each brain-stimulation method are additional factors likely to influence the different response times.

When using transducers directly coupled to the skin of mice, bilateral stimulation with transcranial US produced the near-simultaneous activation of several muscle groups, indicated by tail, forepaw, and whisker movements ( Movie S2 ). By using acoustic collimators having an output aperture of d = 2.0, 3.0, or 4.7 mm and by making small (≈2 mm) adjustments to the positioning of transducers or collimators over motor cortex within a subject, we could differentially evoke the activity of isolated muscle groups ( Movie S2 ). Despite these intriguing observations, we found it difficult to reliably generate fine maps of mouse motor cortex using US for brain stimulation. The likeliest explanation for this difficulty is that the topographical/spatial segregation of different motor areas represented on the mouse cortex are below the resolution limits of US (see Spatial Resolution of Brain Circuit Activation with Transcranial Pulsed Ultrasound below).

We next acquired fine-wire electromyograms (EMG) and videos of muscle contractions in response to US stimulation of motor cortex in skin- and skull-intact, anesthetized mice ( Movie S1 ). Using transcranial US to stimulate motor cortex, we evoked muscle contraction and movements in 92% of the mice tested. The muscle activity triggered by US stimulation of motor cortex produced EMG responses similar to those acquired during spontaneous muscle twitches ( Figure 3 A ).

(E) Raw (black) and averaged (gray; ten trials) temperature recordings obtained from motor cortex in response to transmission of US waveforms with short pulse durations (PD) used in stimulus waveforms (top). Similarly, temperature recordings of cortex in response to waveforms having a PD ∼100 times longer than those used in stimulus waveforms (middle and bottom).

(C) EMG failure probability histograms are shown for four progressively increasing stimulus repetition frequencies (left; Movie S3 ). Raw US-evoked EMG traces are shown for two different stimulus repetition frequencies (right). Data shown are mean ± SEM.

(B) EMG response latencies (top) and amplitudes (bottom) recorded from the left triceps brachii in response to right motor cortex stimulation are plotted as a function of trial number repeated at 0.1 Hz. Individual US-evoked raw EMG traces are shown for different trials (right).

(A) Raw (left) and full-wave rectified (FWR; right) EMG traces obtained for a spontaneous muscle twitch (top) and average (ten trials) increase in muscle activity produced by transcranial US stimulation of motor cortex (bottom; Movie S1 ). The duration of the US stimulus waveform (black), average US-evoked EMG trace (gray), and EMG integral (gray dashed line) are shown superimposed at lower right.

We first studied the influence of pulsed US on intact motor cortex because it enables electrophysiological and behavioral measures of brain activation ( Movie S1 ). We recorded local field potentials (LFP) and multiunit activity (MUA) in primary motor cortex (M1) while transmitting pulsed US (0.35 MHz, 80 c/p, 1.5 kHz PRF, 100 pulses) having an I= 36.20 mW/cmthrough acoustic collimators (d = 4.7 mm) to the recording locations in anesthetized mice (n = 8; Figures 2 A and 2B ). Pulsed US triggered an LFP in M1 with a mean amplitude of −350.59 ± 43.34 μV ( Figure 2 B, 25 trials each). The LFP was associated with an increase in the frequency of cortical spikes ( Figures 2 C and 2D). This increase in spiking evoked by pulsed US was temporally precise and apparent within 50 ms of stimulus onset ( Figure 2 D). We found a broad range of pulsed US waveforms were equally capable of stimulating intact brain circuits as discussed below. Application of TTX (100 μM) to M1 (n = 4 mice) attenuated US-evoked increases in cortical activity, indicating that transcranial US stimulates neuronal activity mediated by action potentials ( Figure 2 B). These data provide evidence that pulsed US can be used to directly stimulate neuronal activity and action potentials in intact brain circuits.

(D) A poststimulus time histogram illustrates the average MUA spike count recorded 500 ms prior to and 500 ms following the delivery of US stimulus waveforms to motor cortex. Data shown are mean ± SEM.

(C) The spike raster plot illustrates the increase of cortical spiking as a function of time in response to 25 consecutive US stimulation trials.

(B) (Top) Raw (black) and average (gray; 25 trials) US-evoked MUA recorded from M1 cortex in response to the delivery of pulsed US waveforms ( Movie S1 ). (Middle) Addition of TTX to the cortex reduced synaptic noise and attenuated US-evoked MUA. (Bottom) Raw control (black), average control (green), and average TTX (red) LFP recorded from M1 cortex in response to 25 US stimulus waveforms delivered every 10 s.

Single US pulses contained between 80 and 225 acoustic cycles per pulse (c/p) for pulse durations (PD) lasting 0.16–0.57 ms. Single US Pulses were repeated at pulse repetition frequencies (PRF) ranging from 1.2 to 3.0 kHz to produce spatial-peak temporal-average intensities (I) of 21–163 mW/cmfor total stimulus duration ranging between 26 and 333 ms. Pulsed US waveforms had peak rarefactional pressures (p) of 0.070–0.097 MPa, pulse intensity integrals (PII) of 0.017–0.095 mJ/cm, and spatial-peak pulse-average intensities (I) of 0.075–0.229 W/cm Figures 1 A, 1B, S1, and S2 illustrate the strategy developed for stimulating intact brain circuits with transcranial pulsed US. The attenuation of US due to propagation through the hair, skin, skull, and dura of mice was <10% ( Figure 1 C), and all intensity values reported were calculated from US pressure measurements acquired using a calibrated hydrophone positioned with a micromanipulator inside fresh ex vivo mouse heads at locations corresponding to the brain circuit being targeted.

We constructed US stimulus waveforms and transmitted them into the intact brains of anesthetized mice (n = 192; Figure 1 A ). The optimal gains between transcranial transmission and brain absorption occurs for US at acoustic frequencies (f) ≤ 0.65 MHz (). Thus, we constructed transcranial stimulus waveforms with US having f = 0.25–0.50 MHz. Intensity characteristics of US stimulus waveforms were calculated based on industry standards and published equations developed by the American Institute of Ultrasound Medicine, the National Electronics Manufacturers Association, and the United Stated Food and Drug Administration (; see Experimental Procedures ).

(C) Projected from a transducer surface to the face of a calibrated hydrophone, the acoustic pressure generated by a 100 cycle pulse of 0.5 MHz ultrasound is shown (left). The pressure generated by the same US pulse when transmitted from the face of the transducer through a fresh ex vivo mouse head to regions corresponding to motor cortex (0.8 mm deep) is shown (right).

(B) An example low-intensity US stimulus waveform is illustrated to highlight the parameters used in their construction. The acoustic intensities generated by the illustrated stimulus waveform are shown in the yellow box.

(A) Illustration of the method used to construct and transmit pulsed US waveforms into the intact mouse brain. Two function generators were connected in series and used to construct stimulus waveforms. An RF amplifier was then used to provide final voltages to US transducers (see Figures S1 and S2 and Experimental Procedures ).

Discussion

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Price L.R. Focused ultrasound modifications of neural circuit activity in a mammalian brain. Mihran et al., 1990 Mihran R.T.

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Huang C.C. In vitro effects of ultrasound with different energies on the conduction properties of neural tissue. To date, previous studies detailing the effects of US on neuronal activity have fallen short of providing methods for its practical implementation in stimulating intact brain function. Prior studies examined the effects of US on neuronal activity by presonicating nervous tissues with US before examining its consequence on electrically evoked activity. These studies indeed revealed how US differentially affects the amplitude and duration of compound action potentials/field potentials evoked with traditional stimulating electrodes (). In other words, previous studies showed that US is capable of modulating electrically evoked activity but not that it alone could stimulate neuronal activity. We have provided clear evidence that transcranial pulsed US can stimulate intact brain circuits without requiring exogenous factors or surgery.

Huerta and Volpe, 2009 Huerta P.T.

Volpe B.T. Transcranial magnetic stimulation, synaptic plasticity and network oscillations. Angel and Gratton, 1982 Angel A.

Gratton D.A. The effect of anaesthetic agents on cerebral cortical responses in the rat. Goss-Sampson and Kriss, 1991 Goss-Sampson M.A.

Kriss A. Effects of pentobarbital and ketamine-xylazine anaesthesia on somatosensory, brainstem auditory and peripheral sensory-motor responses in the rat. Our observations also serve as preliminary evidence that pulsed US can be used to probe intrinsic characteristics of brain circuits. For example, US stimulation of motor cortex produced short bursts of activity (<100 ms) and peripheral muscle contractions, whereas stimulation of the hippocampus with similar waveforms triggered characteristic rhythmic bursting (recurrent activity), which lasted 2–3 s. These observations lead us to question whether stimulation of a given brain region with US can mediate even broader circuit activation based on functional connectivity. Such abilities have been shown and discussed for other transcranial brain-stimulation approaches like TMS (). Future studies should be designed to study the influence of US on activity in corticothalamic, corticocortical, and thalamocortical pathways as we have done here for corticospinal circuits. Similar to widely recognized observations using other cortical-stimulation methods (), we found that the success of brain activation with transcranial pulsed US was dependent on the plane of anesthesia. When mice were in moderate to light anesthesia planes (mild responsiveness to tail pinch), we found that US-evoked activity was highly consistent across multiple repeated trials as described above.

Ang et al., 2006 Ang Jr., E.S.

Gluncic V.

Duque A.

Schafer M.E.

Rakic P. Prenatal exposure to ultrasound waves impacts neuronal migration in mice. Although our observations indicate that pulsed US provides a safe mode of brain stimulation in mice ( Figure 6 and S4 ), it should not be inferred that the same is true for other animal species. Safety studies in other animals are required for any such conclusions to be drawn. Since we suspect that standing waves may inadvertently influence the activity of some brain regions under certain conditions, future studies should attend to the influence of such reflections on brain tissue, regardless of the focusing method implemented. This is particularly true for cases where high-intensity ultrasound may be used to treat brain tissues as discussed below. The less-direct safety implications of our study also need to be considered. Diagnostic fetal US has been shown to disrupt neuronal migration in developing rat fetal brains (). Those effects could be due to the influence of US on neuronal activity or growth factor expression patterns in developing fetal brains. Having dire ramifications on the global use of diagnostic fetal ultrasound, investigations into such possibilities are warranted.

Hynynen et al., 2004 Hynynen K.

Clement G.T.

McDannold N.

Vykhodtseva N.

King R.

White P.J.

Vitek S.

Jolesz F.A. 500-element ultrasound phased array system for noninvasive focal surgery of the brain: a preliminary rabbit study with ex vivo human skulls. Hynynen et al., 2006 Hynynen K.

McDannold N.

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Zadicario E.

Killiany R.

Moore T.

Rosen D. Pre-clinical testing of a phased array ultrasound system for MRI-guided noninvasive surgery of the brain—a primate study. Martin et al., 2009 Martin E.

Jeanmonod D.

Morel A.

Zadicario E.

Werner B. High-intensity focused ultrasound for noninvasive functional neurosurgery. 2) to perform noninvasive thalamotomies (d = 4.0 mm) for the treatment of chronic neuropathic pain by focusing US through the intact human skull to deep thalamic nuclei using phased arrays ( Martin et al., 2009 Martin E.

Jeanmonod D.

Morel A.

Zadicario E.

Werner B. High-intensity focused ultrasound for noninvasive functional neurosurgery. Focusing of US through skull bones, including those of humans, can be achieved using transducers arranged in phased arrays (). A recent clinical study reported using transcranial MRI-guided high-intensity focused ultrasound (0.65 MHz, >1000 W/cm) to perform noninvasive thalamotomies (d = 4.0 mm) for the treatment of chronic neuropathic pain by focusing US through the intact human skull to deep thalamic nuclei using phased arrays (). These abilities to focus US through the intact skull into the deep-brain regions certainly raise the possibility of using pulsed US in the noninvasive stimulation of human brain circuits. However, cautiously conducted preclinical safety and efficacy studies are required across independent groups before it can be determined if pulsed US might be useful in such an application.