NeuroImage 124 (2016) 752–761

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Transcranial functional ultrasound imaging of the brain usingmicrobubble-enhanced ultrasensitive Doppler

Claudia Errico a,b,c, Bruno-Félix Osmanski a,b,c, Sophie Pezet d,e, Olivier Couture a,b,c,Zsolt Lenkei d,e,1, Mickael Tanter a,b,c,⁎,1a INSERM, Institut Langevin, 1 rue Jussieu, 75005, Paris, Franceb ESPCI ParisTech, PSL Research University, Institut Langevin, 1 rue Jussieu, 75005, Paris, Francec CNRS, Institut Langevin, 1 rue Jussieu, 75005, Paris, Franced CNRS, UMR 8249, 10 rue Vauquelin, 75005 Paris, Francee Brain Plasticity Unit, ESPCI-ParisTech, PSL Research University 10 rue Vauquelin, 75005 Paris, France

⁎ Corresponding author at: Institut Langevin, ESPCI-PariFrance.

E-mail address: Mickael.Tanter@espci.fr (M. Tanter).1 Contributed equally, ZL and MT are co-last authors.

http://dx.doi.org/10.1016/j.neuroimage.2015.09.0371053-8119/© 2015 The Authors. Published by Elsevier Inc

a b s t r a c t

a r t i c l e i n f o

Article history:Received 30 April 2015Accepted 14 September 2015Available online 28 September 2015

Keywords:TranscranialFunctional ultrasound imagingMicrobubblesPrimary sensory cortexBlood volumeSomatosensory activation

Functional ultrasound (fUS) is a novel neuroimaging technique, based on high-sensitivity ultrafast Doppler imag-ing of cerebral blood volume, capable ofmeasuring brain activation and connectivity in rodents with high spatio-temporal resolution (100 μm, 1 ms). However, the skull attenuates acoustic waves, so fUS in rats currentlyrequires craniotomy or a thinned-skull window. Here we propose a non-invasive approach by enhancing thefUS signal with a contrast agent, inert gas microbubbles. Plane-wave illumination of the brain at high framerate (500 Hz compounded sequence with three tilted plane waves, PRF = 1500Hz with a 128 element 15 MHzlinear transducer), yields highly-resolved neurovascular maps. We compared fUS imaging performance throughthe intact skull bone (transcranial fUS) versus a thinned-skullwindow in the same animal. First, we show that thevascular network of the adult rat brain can be imaged transcranially only after a bolus intravenous injection ofmicrobubbles, which leads to a 9 dB gain in the contrast-to-tissue ratio. Next, we demonstrate that functional in-crease in the blood volume of the primary sensory cortex after targeted electrical-evoked stimulations of the sci-atic nerve is observable transcranially in presence of contrast agents, with high reproducibility (Pearson'scoefficient ρ = 0.7 ± 0.1, p = 0.85). Our work demonstrates that the combination of ultrafast Doppler imagingand injection of contrast agent allows non-invasive functional brain imaging through the intact skull bone in rats.These results should ease non-invasive longitudinal studies in rodents and open a promising perspective for theadoption of highly resolved fUS approaches for the adult human brain.

© 2015 The Authors. Published by Elsevier Inc. This is an open access article under the CC BY-NC-ND license(http://creativecommons.org/licenses/by-nc-nd/4.0/).

Introduction

The functional activity of the brain can be followed through mea-surement of its blood supply, as first proposed by Lavoisier (1920) anddemonstrated by Mosso (1881). Today, several functional imaging mo-dalities exploit, as their experimental read-out, local hyperemia to mapthe functional response of the brain to stimuli. These include functionalmagnetic resonance imaging (fMRI) (Ogawa et al., 1990; Kim andUgurbil, 2003; Rossi, 2006; Richiardi et al., 2011), and other functionalimaging techniques such as optical coherence tomography and multi-photon microscopy (Sheth et al., 2003; Chen et al., 2009). Each one of

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these techniques has its own specificity, sensitivity and spatiotemporalresolution, and have different limitations. These restrictions currentlyhinder the broad dissemination of functional imaging approaches intranslational and pre-clinical research settings. Therefore, the validationof complementary or alternativemethods for in vivo imaging of local hy-peremia in rodents, a major pre-clinical animal model, is an importantscientific objective.

Recently, very high frame rate ultrasound imaging (10,000 framesper second) (Tanter and Fink, 2014) was shown to enable high-resolution and high-sensitivity power Doppler imaging (Bercoff et al.,2011). Applied to brain imaging, it led to high sensitivitymapping of ce-rebral blood volume (CBV) and functional ultrasound (fUS) imaging oftask-evoked changes in cortical activity in the rat brain (Macé et al.,2011; Mace et al., 2013). fUS is able to detect subtle changes of lowblood flow in small cerebral vessels with a very high sensitivity. It con-sists in replacing the conventional line-by-line scanning of tissue withfocused beams by successive transmissions of either an ultrasonic

the CC BY-NC-ND license (http://creativecommons.org/licenses/by-nc-nd/4.0/).

753C. Errico et al. / NeuroImage 124 (2016) 752–761

plane or diverging waves, to attain frame rates close to 20,000 framesper second at 3 cm depth (Osmanski et al., 2012b). At these frame-rates, fUS can differentiate slow-moving blood (down to 1 mm/s),from surrounding tissue (Mace et al., 2013), and consequently, it is po-tentially able to detect blood flow changes in the capillary network evenif its spatial resolution (~100 μm) is not sufficient to individually mapthese vessels. Such sensitivity is required to detect slight variations inblood flow following sensory stimulation, since red blood cell velocitiesin small cortical arterioles and capillaries are typically in the 1–10mm/srange (Schaffer et al., 2006; Shih et al., 2009; Kobat et al., 2011; Shihet al., 2012). This elevated sensitivity allowed fUS to track the spatio-temporal dynamics of the activation within the rat brain followingwhisker stimulation or during induced epileptic seizures (Macé et al.,2011). It was also shown to be sensitive enough to detect odor-evoked stimulation (Osmanski et al., 2014a) and, remarkably, to mapfunctional connectivity in the living rat brainwith amuchhigher spatio-temporal resolution than fMRI (Osmanski et al., 2014b).

However, albeit ultrasound can propagate deep within tissues,cerebral fUS remains limited by its poor penetration through theskull. An attenuation of 6.9 dB/cm at 15 MHz has been observed inprevious studies (Fry and Barger, 1978; Goss et al., 1979; Larratet al., 2010; Pinton et al., 2012), which renders small variations inblood flow indistinguishable from noise. Therefore, in order tostudy changes in functional blood volume in microvessels deep inthe rat brain both elevated resolution (typically 100 μm resolutionin-plane) and penetration (up to 20 mm penetration depth) is re-quired. Surgical procedures involving a craniotomy (Mace et al.,2013) or a bilateral thinned-skull window (Osmanski et al., 2014b)meet such requirements, but their invasiveness restrains large dis-semination of this neuroimaging technique. Similarly, clinical appli-cation of fUS is limited today to patients with a pre-existing cranialwindow such as the fontanel in newborns (Demene et al., 2014).

An alternative to surgical enhancement of penetration depthwould be to enhance the fUS signal strength. Importantly,clinically-available solutions of injectable microbubbles can beused to increase the signal from blood vessels (Frinking et al.,2000). Because of the high impedance mismatch between the sur-rounding tissue and the gas content of the microbubbles, further en-hanced by their acoustically nonlinear behavior (Simpson et al.,1999; Goertz et al., 2005), ultrasound scattering is increased by upto 35 dB (Burns and Wilson, 2006). These agents have been used toincrease Doppler signal in the myocardium (Klibanov, 2013), in theliver, in kidneys and in tumors (Williams et al., 2011), and even inthe rat brain after craniotomy for low-resolution line-by-line scan-ning (van Raaij et al., 2011). Notably, recent studies demonstratedthat microbubble imaging could specially benefit from ultrafastframe rates. For instance, plane-wave imaging reduces microbubbledisruption (Couture et al., 2012) and allows the observation of thedissolution of these agents (Couture et al., 2009). Ultrafast imagingof contrast agents even allowed superlocalization ultrasound imag-ing in sub-wavelength width microfluidic channels (Desailly et al.,2013). These results suggest that combining microbubble injectionwith ultrafast imaging could lead to an important leap in the sensi-tivity of brain fUS, which may even fully compensate for the attenu-ation from the skull, by combining both preserved resolution andincreased penetration depth. This currently unexplored field opensto a noninvasive approach to measure local changes in cerebralblood volume (CBV) of the whole rat brain, through the intactskull. Here we used boluses of intravenously injected microbubblecontrast agents to specifically enhance backscattered vascular sig-nals in the investigated coronal imaging plane. As functional read-outs we first tested the capacity of microbubble contrast agents toobtain a vascular map though the intact skull, then we measuredtask-induced functional hyperemia. Our results show that thatmicrobubble contrast agents allow highly sensitive and non-invasive fUS brain imaging through the intact skull bone in rats.

Material and methods

Animals

All experiments were performed in agreement with the EuropeanCommunity Council Directive of 22nd September 2010 (010/63/UE)and the local ethics committee (Comité d’éthique en matièred’expérimentation animale n°59, C2EA –59, ‘Paris Centre et Sud’, projectagreement # 2014-14). Experiments were performed on male SpragueDawley rats (n = 7, Janvier Labs; Le Genest St Isle, France), weighing200–225 g at the beginning of the experiments. Animals arrived in thelaboratory 1 week before the experiment. They were kept three percage at a constant temperature of 22 °C, with a 12 h alternating light/dark cycle (light 7 AM–7 PM). Food andwater were available ad libitum.

Preparation of bilateral and unilateral thinned-skull imaging windows

To perform ultrasound imaging through the thinned- skull windowof adult rats, the skullwas thinned to 75–100 μmover an area of approx-imately 0.6 cm × 0.9 cm (bilateral window) or of 0.6 cm × 0.4 cm(unilateral window) at 1–2 days before imaging. Under anesthesia (in-traperitoneal injection ofmedetomidine (Domitor®, 0.3mg· kg−1) andketamine (Imalgène®, 40 mg · kg−1), the head of the animal wasplaced in a stereotaxic frame, and the three layers of bone of the skullwere consecutively removed by drilling (Foredom, USA) at low speed,by using a micro drill steel burr (Fine Science Tools, cat. no. 19007-07). During the thinning procedure, the skull was frequently cooledwith saline and an airstream, as described in (Yang et al., 2010),resulting in a lack of heating, swelling, or edema of the cerebral cortex.The thinned-skull windowwas protected by a small (1 cm×1 cm) plas-tic cover, and the skin was sutured using 5.0 non-absorbable Ethiconthread. Preliminary experiments showed that this method enabledgood quality ultrasound imaging for as long as 1 week after skullthinning.

Preparation of microbubble ultrasound contrast agents

Perfluorcarbon-filled microbubbles (Bracco, Plan-Les-Ouates,Switzerland)with a diameter ranging between 1 μmand 5 μmwere dis-solved with 0.9% normal saline solution to a concentration of 2 × 108

microbubbles/ml. Two initial 200 μl boluses of the contrast agent wereinjected in the catheterized jugular vein before starting the hind limbprimary sensory cortex (S1HL) activations to ensure that the fluctua-tions of baseline values were less than 5%. Consecutively, we delivered,every 10 min, a maximum of 11–13 bolus injections (corresponding to2.6 ml of the initial suspension). Preliminary experiments establishedthat the concentration of microbubbles was stable 1 min after each in-jection, time at which the functional ultrasound acquisition was initial-ized. Hence, the dynamic time-course changes of the bolus injectionwere investigated by imaging for 10 min the temporal evolution of thecirculating bubbles in the blood stream.

fUS imaging

Animalswere anesthetized using an initial intraperitoneal (IP) injectionof medetomidine (Domitor®, 0.3 mg · kg−1) and ketamine (Imalgène®,40 mg · kg−1), followed by hourly IP injections of medetomidine(0.1 mg · kg−1 h−1) and ketamine (12.5 mg · kg−1 h−1) in order tokeep the animals under stable anesthetized conditions. The jugularvein was catheterized in order to perform the injections of microbubblecontrast agents. The thinned skull was rinsed with sterile saline and1 cm3 of ultrasound coupling gel was placed on the window. The ani-mals were then placed in a stereotaxic frame (Stoelting; Chicago, IL,USA).

A 15 MHz linear transducer array (128 elements, 8 mm elevationfocus, 80 μm pitch) was directly installed on the head of the animal,

754 C. Errico et al. / NeuroImage 124 (2016) 752–761

connected to the ultrafast research ultrasound scanner (AixPlorer, Su-perSonic Imagine, Aix-en-Provence, France) and coupled on the brainvia standard degased ultrasound gel. Real time B-mode imaging wasinitially used to control the placement of the probe on the field ofview. The software-based architecture of the scanner enabled Matlab(MathWorks; Natick, Massachusetts, USA) programming of customtransmit/receive ultrasound sequences. Each session began with ananterior-posterior Doppler scan, used to visualize the shape of theblood vessels, and identify the stereotaxic coordinates. fUS acquisitionswere achieved at the level of the choroid plexus of the lateral ventricle,corresponding to the anteroposterior coordinate Bregma −1 mm.

Body temperature was maintained constant using a heating pad(Gaymar Industries; New York, NY, USA). A 1 cm incision was made inthe skin and muscle of both hindpaws, above the femur in order to ex-pose and isolate the sciatic nerve. A hook-shaped stimulating electrodewas gently inserted on the left and right sciatic nerve (LSN and RSN, re-spectively) to alternatively induce the electric sciatic stimulations. Thenerve was allowed to rest for 10 min between each stimulation se-quence. Trains of five C fiber stimulations (25 pulses at 5 Hz, 0.2 mA,and 100 μswidth)were delivered for 5 s on each sciatic nerve, separatedby 20 s OFF period to recover the baseline (Fig. 1A). Our exclusion crite-rion was based on the physiological condition of the animals. Only ani-mals whose fluctuations of the power Doppler signal baseline (absence

Fig. 1. Schematic view of the fUS imaging setup and fUS imaging protocol. A: The 15MHz probeframe immobilized the skull. An isolated constant current stimulator delivered five electrical stimpresents the stimulation parameters. B: The continuous-fUS insonification sequence consisted iblood flowwithout aliasing. The resulting three low quality images were coherently summed ttimes leading to 150 s acquisition time to stack a set of 200 high quality images per each sequentrical-evoked activation of the brain.

of stimulation) were stably b5% were included in the study. Experi-ments were performed on n = 7 male Sprague–Dawley rats, by usingthree different protocols: bilateral window (BW, n = 1, protocol A,Fig. 2A), unilateral thinned-skull window (TSW, n = 4, protocol B,Fig. 2B) or full transcranial (IS, n = 2 protocol C, Fig. 2C) imaging. Thisled respectively to the fUS imaging of n = 12 transcranial and of n =12 thinned skull hemispheres (Fig. 5).

Transcranial ultrafast Doppler imaging sequence

Functional ultrasound (fUS) imaging was performed using UltrafastDoppler (Bercoff et al., 2011) based on compounded plane-wave ultra-sound imaging (Tanter et al., 2002; Montaldo et al., 2009). This tech-nique has sufficiently high sensitivity to detect subtle hemodynamicchanges due to the neurovascular coupling with a high spatio-temporal resolution (1 ms, 100 μm).

First, the vascularization map of the brain was obtained by using adedicated MATLAB script to move the US probe, fixed on a three-axesmicrostep motor, 0.5 mm after each ultrafast Doppler acquisition. Thescanning sequence consisted in 15 tiled planar ultrasonic waves (−14°to 14° angles, with a step of two angles 2°) with a pulse repetition fre-quency (PRF) of 7500 Hz. Once the choroid plexus of the lateral ventriclewas identified, the ultrasound probewasfixed in coronal orientation over

was placed on the top of the animals' heads and coupled via ultrasound gel. A stereotaxiculi alternatively on the left and right sciatic nerves (LSN and RSN, respectively). The panel

n three tilted planewaves (−3°, 0°and 3°). The frame rate was set to 500 Hz to sample theo obtain a single high quality B-mode image of the brain. The sequence was repeated 374ce. The power Doppler images were acquired over 150 s in order to fully recover the elec-

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Fig. 2.High-resolution ultrafast Doppler imaging of the rat brain vasculature through the intact skull (IS) after intravenous injection ofmicrobubble contrast agents.A–C: Schematic of thethree imaging configurations used: bilateral thinned skull window (TSW, A), intact skull (IS, B) or unilateral thinned skull window (C).D–F: Examples of vascular maps obtained with thedifferent imaging modalities in absence (1) or presence of microbubbles (2). D: High quality imaging of both superficial and deep cerebral structures was obtained in the control config-uration, i.e. without microbubbles through the bilateral thinned-skull window. E–F: In absence of contrast agent, the microvascular structures of the brain are no longer distinguishablethrough the intact skull (IS). However, the loss of signal due to the cranial bone is fully compensated by the injection of 150 μl ofmicrobubble contrast agents. The yellowboxes inD–F showthe fUS signal in the choroid plexus of the lateral ventricle, allowing the correct US probe placement above the S1HL at Bregma −1 mm. Scale bar: 2 mm.

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the hind limb primary sensory cortex (S1HL) at the anterior-posterior co-ordinate of Bregma −1 mm. Over the thinned-skull window, the align-ment was performed without the addition of contrast agent. In the fulltranscranial protocol, an initial injection of 150 µl microbubbles was nec-essary to highlight the blood vessels and achieve a correct alignment ofthe probe over the reference coordinates. Activation maps of the S1HLcortex were collected with a continuous-fUS activation sequenceconsisting in a compounded plane wave sequence at 500 Hz frame rateusing three tilted plane waves (−3°,0° and 3°, PRF = 1500Hz).

The backscattered echoes were recorded, beamformed and coher-ently added to produce a stack of high quality B-mode echographic im-ages (Fig. 1B). As the intensity of the power Doppler images isproportional to the cerebral blood volume (CBV), we implemented anincoherent mean of the blood signal to cancel blood pulsatility and ob-tain high quality ultrafast Doppler images. We have previously demon-strated that this incoherent temporalmean is equivalent to a filtering ofpulsatile time component thanks to the very high temporal sampling ofblood flow signals (Osmanski et al., 2014a) We repeated thecontinuous-fUS sequence 374 times, each containing 200 frames, inorder to recover for 150 s the entire duration of the five electrical stim-uli. The electrical stimuli generated by the isolated constant currentstimulator and the ultrafast Doppler sequence were triggered withtwo arbitrary waveform generators.

Treatment of fUS data

The acquisition and reconstruction of both the vascularization and ac-tivation maps were carried out on the programmable US scanner withdedicated MATLAB scripts. Backscattered signals from the brain were fil-tered to distinguish the moving red blood cells and microbubbles fromthe static tissue. Singular value decomposition (SVD) was implementedon the stack of the fUS images and the largest Eigenvalues were eliminat-ed to filter the slowest variations in the Power Doppler signal related tothe tissue signal (Demene et al., 2014). This multidimensional (spaceand time) filter is able to discriminate very low speed blood flows (upto 1mm/s) from tissue displacements by exploiting the difference of spa-tiotemporal coherence between tissue and blood flow signals. Then, thebackscattered signals were filtered with a 4th order Butterworth high-

pass filter with a cut-off frequency of 75 Hz to further remove tissueand motion artifact coming from the skull. The Doppler signal was origi-nally obtained by the incoherent temporal mean of the blood signal. Thecerebral blood volume increase (CBV) evoked by the functional stimula-tions was measured in each pixel, which were 100 × 80 μm in size witha slice thickness of 300 μm.

The activation maps were computed from the temporal correlationbetween the stimuli and the increase in CBV. We observed a contralat-eral activation of the brain in response to the evoked task. The activatedpixelswere considered significant for a correlation ρ N 2σ, where ρ is thespatial correlation of the activation maps and σ their spatial standarddeviation. The Doppler signal was averaged over time in the region ofinterest (corresponding to activated areas) to retrieve step by step thetime course of the entire electrical stimulation pattern. Its amplitudewas represented as a percentage of change relative to the baselines ofthe activated areas ± standard deviation. Furthermore, the hemody-namic signals from the ROI (activated areas) were extracted to estimatethe temporal correlation between the evoked task and themeasured in-crease in blood volume. The reproducibility of the activationwas evalu-ated,with andwithout the presence of contrast agent, using the Pearsoncorrelation coefficient (R).

Results

Ultrafast Doppler scan through the thinned-skull windowor through the in-tact skull

First, we compared three different ultrasound imaging configura-tions: imaging through bilateral and unilateral thinned-skull windows(TSW) and full transcranial imaging through the intact skull (IS), withand without intravenous injection of microbubbles (Fig. 2A, B, C). Asdemonstrated previously (Macé et al., 2011; Osmanski et al., 2014a),the vascular network of the rat cortex could be mapped with ultrafastDoppler imaging through a TSW (Fig. 2D1, F1). Here, the choroid plexusof the lateral ventricle was used as a reference point for the localizationof the imaging plane (Bregma −1 mm), which contains the hind limbprimary sensory cortex (S1HL) (Paxinos and Watson, 2006).

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The intact skull attenuated the high-frequency ultrasonic waves(15 MHz emitting frequency) hindering the acquisition of the vascularmap (Fig. 2E1, F1). Strikingly, after the injection of microbubbles, thisloss of signal was fully compensated and blood vessels as deep as12.5 mm could be visualized (Fig. 2E2). For better visualization of themicrobubble-induced signal amplification, we compared the signal ob-tained in the two hemispheres of a unilateral TSW configuration(Fig. 2F2). A typical example of microbubble-enhanced fUS rostro-caudal scan through the intact skull is presented in Fig. 3, showingthat a large part of the rat brain can be visualized up to 12.5 mmdepth using micro-bubbles as a contrast agent. We have calculated thecontrast-to-tissue ratio (CTR), defined as the ratio of the scatteredpower from the blood echo (combining red blood cells andmicrobubbles) and that from the surrounding tissue. In presence ofmicrobubbles, we calculated a gain of 9 dB in the Power Doppler pre-stimulus baseline prior to normalization. Hence, the use of contrastagent reduces the relative contribution of the noise in the Power Dopp-ler estimates as the fUS signal is generally close to the noise floor of theelectronic system. These results show that by using inert gasmicrobubbles as a contrast agent, significant fUS signal amplificationcan be achieved, allowing mapping of the cerebral vasculature throughthe intact rat skull. The vascularizationmaps obtained through the IS vs.TSWwere comparable (Fig. 2D–F). Hence, in spite of the presence of thebone, we were able to perform functional imaging of the brain withoutusing aberration correction methods.

Microbubble-enhanced imaging of functional hyperemia through the intactskull

The second step of the studywas to investigatewhether contrast en-hancement by microbubble contrast agents also enables the measure-ment of functional hyperemia through the intact skull. The activationmaps shown in Fig. 4 display the spatial distribution of thehemodynam-ic response in the S1HL to the electrical stimulation of the sciatic nerve.The grayscale background image shows the vascular network detectedby ultrafast ultrasound imaging, through the intact skull or a thinned-skull window, with or without an injection of microbubbles in singlerepresentative animals. The overlaid correlationmap shows the tempo-ral correlation between the five stimuli and the evoked change in thePower Doppler (PD) signal. As a reference, we first performed the acti-vation of the sensory cortex on the bilateral TSW protocol without theinjection of contrast agents (Fig. 4A). The set of five stimuli were repeat-ed up to a maximum of three times on the left or right sciatic nerves(LSN and RSN, respectively) on a total of n = 12 IS hemispheres andn = 12 TSW hemispheres (Fig. 5).

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Fig. 3. Antero-posterior fUS scans allow to map the vasculature of the rat brain through the intagent led to a clear ultrasound signal at all different coronal imagingplanes, at full depth. The steDoppler acquisition. Scale bar: 2 mm.

As previously described (Osmanski et al., 2014a), using a thinned-skull window in absence of micro-bubbles, the stimulation of the rightand left sciatic nerve induce an evoked hemodynamic response in thecontralateral area of the S1HL (Fig. 4A2, A3, D2). However, in absenceof microbubbles, such evoked hemodynamic response cannot be ob-served through the intact skull (Fig. 4B1, B2, B3, D3). Strikingly, intrave-nous injection of microbubbles allowed the observation of the brainvasculature through the intact skull (Fig. 4C1), allowing the measure-ment of stimulation-induced hemodynamic changes in the right andleft S1HL (Fig. 4C2, C3, E2, E3). Comparison of the activation maps re-trieved from the thinned-skull hemispheres without injection ofmicrobubbles with those acquired transcranially with the use of con-trast agents showed that transcranial Ultrafast Doppler imaging in com-bination with ultrasound contrast agents could detect reproducibly(Pearson correlation coefficient of 0.7 ± 0.1, p = 0.85) and non-invasively functional changes deep in the rat brain.

Sensitivity and robustness of contrast-agent enhanced transcranial Dopplerimaging

In order to further characterize the performance of contrast agentenhanced transcranial imaging, we observed the time-resolved powerDoppler (PD) signal in the activated area with a high-frame rate imag-ing sequence (500 frames per second), in the TSW and IS configurationswithout and with contrast agent injection. As shown in Fig. 6A, thisallowed the characterization of the temporal evolution of the hemody-namic response to the evoked task, i.e. the set of five stimuli. For eachhemisphere, the intensity of the fUS signal was normalized to the pre-stimulus baseline andwas represented in % changes of the Power Dopp-ler signal. Notably, the PD increase corresponding to the vascular re-sponse was consequently higher after the first stimulation than afterthe following four, corresponding to around 2% diminution in peak-height between the first and the fifth peaks, in the thinned-skull config-uration without contrast agent. This gradual diminution of peak ampli-tude, most likely due to the physiological adaptation to the electricalstimulations, was further enhanced with the use of contrast agents, be-cause of the gradual elimination of the microbubble bolus from the cir-culation (Fig. 6B). We could correct for this second source of signaldiminution by continuously measuring the power Doppler intensity inthe sagittal sinus and by adjusting the temporal evolution of the hemo-dynamic response function by the slope of this reference intensity.The resulting corrected curve (blue on Fig. 6B) was comparable to thecurve obtained with the thinned-skull configuration in absence ofmicrobubbles, where no correctionwas needed. After these corrections,the mean temporal hemodynamic parameters were retrieved by aver-aging the five electrical stimuli to obtain a typical hemodynamic

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act skull (IS) after an intravenous injection of microbubbles. Injection of 150 μl of contrastreotaxic coordinates (antero-posterior distances to Bregma) are reported on the top of each

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Fig. 4.Microbubble contrast agents allow fUS imaging of cortical hyperemia evoked by electrical stimulation of the sciatic nerve, through the intact skull (IS). A–E, The left column showsthe vascularization maps obtained by power Doppler imaging of the investigated coronal plane at Bregma−1 mm in representative animals for each imaging condition. The middle andthe right columns show the hemodynamic responses evoked by stimulation of either the right (RSN) or the left sciatic nerves (LSN). The correlation coefficient shown in the color barrepresents the temporal correlation between the evoked task and the hemodynamic response recorded in the S1HL. Scale bar: 2 mm.

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response profile in correspondence to the evoked tasks (for individualvalues see Table 1). In Fig. 6C, we report the mean values, the standarderrors across the different trials and across subjects for the three

Fig. 5. Summary of successful trials on each animal for the different fUS imaging configu-rations. Functional activations of the S1HLwere performed on n=7 animals. A distinctionbetween transcranial hemispheres and thinned skull hemispheres led to a total of n= 12transcranial S1HL activation through the intact skull when boluses of microbubbles weredelivered in the bloodstream and n=12 activations on the thinned skull windows (n=6without microbubbles and n = 6 with injection of contrast agent). fUS imaging throughthe intact skull in absence of microbubbles was not applicable (NA).

investigated experimental configurations. An unpaired Student’s t-testwas performed aiming to compare the control case (TSW no injectionof contrast agents), with the other case studies in presence ofmicrobubbles. As shown in Fig. 6D, we measured a 4 ± 1%stimulation-induced increase of the peak amplitude (PA) with the con-trol imaging protocol (i.e. TSW without control agent).

Strikingly, for both protocols involving the injection ofmicrobubbles, the peak amplitude displayed a highly significant in-crease, leading to 12 ± 2% (p b 0.0001) increase in PA for the thinned-skull and 9 ± 2% (p b 0.0001) PA increase in the transcranial configura-tion, as compared to the control imaging protocol (unpaired Student's t-test). In the time-domain, the increase in blood volume was stronglycorrelated with the evoked task (thinned hemispheres without bub-bles: ρ = 0.7 ± 0.1, thinned hemispheres with bubbles: ρ = 0.7 ±0.1, transcranial hemispheres with bubbles: ρ = 0.6 ± 0.2). Finally,dynamic parameters of the hemodynamic response, such as thepeak-to-peak and time-to-peak values were roughly equivalent inthe TSW or IS with microbubbles imaging protocols (Table 1).Onset times, assessed as the time needed to reach 10% of themaximum dilation (Hall et al. 2014) were ~2–3 s for each condition(Table 1). Taken together, these results show that microbubblecontrast agents allow robust and highly sensitive dynamic trackingof activation-evoked hemodynamic changes through the intactskull, significantly surpassing in performance the current state-of-the art thinned-skull imaging configuration.

Fig. 6. Sensitivity and robustness of contrast-agent enhanced transcranial Doppler imaging. A: Averaged hemodynamic response function (n= 6, black curve) evaluated on the thinned-skull hemisphereswithout injection ofmicrobubbles (control imaging protocol). The stimulation pattern is shown in red. Due to physiological adaptation to the electrical stimulations, thepeak amplitude displays a decrease of ~2% from the first to the last activation. B: Time evolution of the hemodynamic responses of a representative animal with the transcranial imagingprotocol (blue curve). The red curve shows the fUS signal profile corrected for the signal diminution, due to the dissolution of the injected boluses of contrast agent. C: Mean evoked he-modynamic response function of the three imaging protocols: thinned-skull window (TSW)without (black curve, n= 6) and with (blue curve, n= 6)microbubbles and intact skull (IS)transcranial hemisphereswithmicrobubbles (red curve, n= 12). The fUS signal in response to the neuronal stimulation is represented as themean± standard deviation across trials andacross subjects. D: Peak enhancement of the three imaging protocols: TSW hemispheres with microbubbles (blue, n = 6); 2, transcranial (IS) hemispheres with microbubbles (red, n =12); 3, TSW hemispheres without microbubbles (black, n = 6). Statistics: paired Student's t-test, compared to the control condition (TSW no bubbles). P values: ****P b 0.0001.

758 C. Errico et al. / NeuroImage 124 (2016) 752–761

Discussion

In the past few years, functional ultrasound (fUS) has been validatedas a novel neuroimaging modality in rodents. However, its applicationhave been limited by the presence of the skull which attenuates the ul-trasonic waves. The objective of our work was to combine the high spa-tiotemporal resolution and sensitivity of fUS with contrast-agentinduced signal enhancement for transcranial applicability. Our resultsshow that after intravenous microbubble injection, vascular structureand hemodynamics can be imaged with excellent spatiotemporal reso-lution through the intact rat skull. Consequently, the vascularmap couldbe obtained over several parallel slices and the choroid plexus of the lat-eral ventricles could be identified as an orientation landmark.Microbubbles-aided transcranial fUS was able to detect the variationof the power Doppler intensity in the primary sensory cortex (S1HL),in strong temporal correlation with the stimulation of the sciaticnerve. Therefore, in presence of contrast agents, the sensitivity of ultra-fast Doppler can greatly benefit from the strong enhancement of rawul-trasonic signals (gain of 9 dB in the contrast-to-tissue ratio),overcoming the attenuation of ultrasonic waves by the cranium. We

Table 1Hemodynamic parameters of the temporal evolution of the power Doppler signal in responsecontrol condition (thinned-skull hemispheres, no bubbles).

Thinned-skull (TSW) hemispheresWith bubbles

Peak-to-peak (s) 25.0 ± 0.3Peak amplitude (%) 12 ± 2 (p b 0.0001)Onset time (s) 2.5 ± 0.8 (p b0.2)Time-to-peak (s) 6.0 ± 1.0 (p b0.07)

have already successfully imaged the vascular dynamics of the ratbrain with an ultrafast Doppler acquisition on craniotomized animalswithout contrast agents by delivering whisker stimuli (Macé et al.,2011) as well as odor evoked ones (Osmanski et al., 2014a). Comparedto the sciatic nerve stimulation used in the present study, these stimula-tions are more relevant physiologically; however the detection of bloodflow changeswas facilitated by the absence of the skull bone. In the cur-rent study, the injection of microbubbles drastically improved the SNRby completely separating the fUS signal from the noise floor. Hence,we foresee this technique to be sensitive enough to measure the CBVchanges in the brain also after physiologically evoked tasks.

In our study, similar or better results could be attained transcraniallythan by using the previous TSW protocol, after injecting a bolus ofmicrobubbles that increase the intensity of the blood signalwith respectto the tissue. These results show that the signal enhancement created bycontrast agents could compensate the attenuation induced by the skull.In our work, we noticed that, at 15 MHz imaging frequency, no aberra-tion correctionwas needed neither on the beamforming process, nor onthe Doppler images. Indeed, as it can be noticed in Fig. 2(D–E–F), thevascularization maps relative to the intact skull vs. thinned-skull

to the stimuli delivered on the S1HL. Statistics: unpaired Student’s t-test, compared to the

Transcranial (IS) hemispheres Thinned-skull (TSW) hemispheresNo bubbles

25.0 ± 0.3 25.0 ± 0.59 ± 2 (p b 0.0001) 4 ± 12.0 ± 0.6 2.0 ± 0.36.0 ± 0.2 (p b 0.003) 5.0 ± 0.7

759C. Errico et al. / NeuroImage 124 (2016) 752–761

window can be comparable. However, if strong phase distortionswouldhave appeared on our Doppler images, we could have implemented analready existing aberration correction method (Osmanski et al., 2012a)by directly using the information contained in the Doppler images.

Understanding which physiological parameter is measured by thefUS sequence is crucial for the interpretation of functional imagingdata. Power Doppler imaging without contrast agents injections isknown to be directly proportional to the number of moving red bloodcells in the sample volume, i.e. to the local blood volume (Shung et al.,1976; Rubin et al., 1995; Rubin et al., 1997), provided that backscatter-ing properties do no vary over time. Without contrast agents, this as-sumption is valid as backscattering properties of red blood cells, suchas hematocrit and shear rate, remain time-invariant, which is highlyprobable (Shung et al., 1992; Cloutier and Qin, 1997). In microbubble-enhanced fUS, such direct proportionality of the power Doppler signalwith the CBV is compromised, since the backscattering property of theblood depends of the temporal variation of contrast agents concentra-tion. In order to ensure a direct proportionality of the fUS signal withCBV during the experiment, we corrected for the transient decrease ofthe backscattering signal, due to the elimination of the bolus, by usingthe power Doppler intensity within the large sinus vessels as referencevalue to normalize the measured functional signals. This correction oftime-varying backscattering properties enabled compensation for theprogressive decrease of the estimated stimuli responses. After normali-zation, the remaining decrease in the response between the five succes-sive stimulations was comparable to that obtained without contrastagent injection, interpreted as a physiological adaptation to the stimuli.Thus, themeasurement of the hemodynamic responsewith fUS appearsrobust even though microbubbles introduce an important variability inthe powerDoppler intensity curve (as shown in Fig. 6C, the standard de-viation of the normalized curve in the thinned hemisphere configura-tion, is three times higher in presence of micro-bubbles than in thetwo other protocols). The high fluctuations in the hemodynamic re-sponse function (HRF) curves in presence of microbubbles are ex-plained by fluctuations of the backscattering properties as the numberof bubbles recruited in each voxel is stochastically fluctuating in the re-gion of interest as compared to the red blood cells that are continuouslyperfusing the vessels. The spatial pattern of the contralateral transcrani-al activations of the sensory cortex detected with microbubbles wassimilar to the pattern obtained during control activations through thethinned-skull window, without contrast agents. Repeated stimulationsgave highly reproducible responses. However, it has been demonstratedthat the effect of the anesthesia could interfere with the neuronal cou-pling (Masamoto and Kanno, 2012). Hence in the present study we ex-cluded animals whose fluctuation of the power Doppler baseline(without stimulation) was higher than 5%, putatively indicating unsta-ble physiological conditions.

Brief stimuli are ideal to investigate the central layers of the somato-sensory cortex, which is highly filled with capillaries, however theywould not be long enough to achieve vascular delayed compliance(Siegel et al., 2003). Most parameters obtained from the hemodynamicresponse, namely the peak-to-peak, the onset time and the time-to-peak were equivalent whenmeasured through a thinned-skull windowand through the skull withmicrobubbles and are consistent with previ-ous reports (van Raaij et al., 2011) and other imaging modalities (Silvaet al., 2007; Tian et al., 2010; Goloshevsky et al., 2011; Hirano et al.,2011; Yu et al., 2012). Very recently, in vivo 2-photon microscopy di-rectly showed that capillary vasodilatation is not a passive response toarteriole dilation but on the contrary, neural activity first dilate capil-laries by actively relaxing pericytes, temporally preceding arteriole dila-tion (Hall et al., 2014). In that study, during a 15-s long whisker-padstimuli the onset time was ~2.7 s for the capillary response (averagevessel diameter d = 4.4 μm), whereas it was ~3.7 s in the penetratingarterioles (average d = 12.8 μm). In our imaging conditions, measuredonset times were relatively short, falling between 2 and 3 s (Table 1),suggesting that the hyperemic signal measured by fUS is dominated

by the capillary response. In conclusion, we demonstrate highly sensi-tive functional ultrasound imaging through the intact skull of adultrats after intravenous injection of microbubble contrast agents. Hence,this novel technique combines the high spatial and temporal resolutionof fUS, with minimal invasiveness. The use of contrast agents for fUSshould allow easier longitudinal studies in animals and paves the wayto clinical transcranial f-ultrasound of the adult human brain. Notwith-standing, the translation to a clinical application of transcranial imagingof the human brain vasculature will be more challenging as the thick-ness of the skull bone is in the range of severalmillimeters. Thus, we ex-pect an attenuationmuchhigher signal attenuation compared to a smallanimal as well as an important phase distortion of the ultrasonic wavespassing through the cranium (Pinton et al., 2012). To apply transcranialfUS on patients, lower frequencies should be preferred (1 MHz), whichwould reduce resolution, butwould exploit microbubbles echo closer totheir resonance frequency range (1–5 MHz).

Conclusion

Neuro-imaging systems are invaluable tools in the understanding ofthe brain both for fundamental research and clinical diagnosis. Our re-sults show how functional ultrasound (fUS) could become a novel neu-roimaging technique to perform high resolution transcranial imaging ofthe brain. Until today, invasive surgerywas necessary to create anatom-ic acoustic windows in the skull to image the microvasculature and itsfunctional response. Here, by combining the sensitivity of ultrafastDoppler imaging with microbubbles to detect blood flow under theskull, stimuli-induced changes in the cerebral blood volume are evaluat-ed deepwithin the cortex. It leads to non-invasive functional imaging ofthe whole brain at typical 100 μm resolution, facilitating longitudinalstudies and paving the way for future clinical applications.

Acknowledgments

The authors thank Fondation Pierre-Gilles de Gennes for funding ofC. Errico, Bracco Research SA for providing the contrast agent andAgence Nationale de la Recherche (MUSLI) for financial support. Thiswork was supported by LABEX WIFI (Laboratory of Excellence ANR-10-LABX-24) within the French Program “Investments for the Future”under reference ANR-10-IDEX-0001-02 PSL* . The research leading tothese results has also received funding from the European ResearchCouncil under the European Union's Seventh Framework Programme(FP7/2007-2013)/ERC grant agreement no. 339244-FUSIMAGINE.

Appendix A. Processing of ultrasonic signals

Computing ultrasound images

Theuse of ultrasound imaging array allows the transmission of ultra-sonic waves and the reception of the backscattered echoes on the sametransducer. In our study, as reported in the Material and Methods, weused a custom-built array with 128 elements and a central frequencyof 20.3 MHz (pitch = 0.08 mm, elevation focus = 10 mm). Its15.4 MHz bandwidth allowed the use of this probe at a frequency of15 MHz. Once the medium was insonified, the ultrasonic echoes wererecollected and beamformed to obtain a complex ultrasonic images(x, z) using themethod described inMontaldo (2009), x and z refer re-spectively to the lateral position and depth of a spatial pixel. This oper-ation was repeated every 2 ms (we used a 500-Hz frame rate) and weobtained a 3D matrix s(x, z, t) where t described the time at which theimage was computed.

Extracting ultrasound power Doppler signal

The principle of Doppler ultrasound imaging relies on the detectionof the moving red blood cells. In fact, this stack of ultrasound images is

760 C. Errico et al. / NeuroImage 124 (2016) 752–761

the summation of the backscatter echoes recorded from the red bloodcells as well as those coming from the tissue. In our study, we willhave to consider also the presence of contrast agents. Ultrasound im-ages s(x, z, t) an be decomposed as:

s x; z; tð Þ ¼ stissue x; z; tð Þ þ sRBC x; z; tð Þ þ smicrobubbles x; z; tð Þ

where stissue(x, z, t), sRBC(x, z, t) and smicrobubbles(x, z, t) are respectively thesignals coming from the tissue, the red blood cells and the microbubbles.

We also would like to point out that ultrasound contrast agents asmicro bubbles are blood pool agents and they can only be distinguishedfrom the red blood cells and tissue if non-linear imaging is performed. Inour study, we preferred fundamental imaging to harmonic imaging,since the transmitting ultrasonic frequency at 15 MHz was far fromthe resonance frequency range of the microbubbles contrast agents,and thus we would not have been able to benefit from their nonlinearproperties. The signal backscattered from the blood sblood(x, z, t) canthen be written as:

sblood x; z; tð Þ ¼ sRBC x; z; tð Þ þ smicrobubbles x; z; tð Þ:

In order to separate the blood signals from the tissue, we imple-mented a spatiotemporal filter based on the singular value decomposi-tion (SVD) of the stack of the f-US images. Such method is described indetails by Demene et al. (2015).

Subsequently, the ultrasound signal s(x, z, t) can be decomposedalongwith its eigenvalues λi , its spatial eigenvectors ui(x, z) and its tem-poral eigenvectors Ti(t). s(x, z, t) can be written as

s x; z; tð Þ ¼Xi

λiui x; zð ÞTi tð Þ

It is also demonstrated in Demene et al. (2015) that the tissue signalis contained in the first elements of this sum. sblood(x, z, t) can be obtain-ed by withdrawing the ntissue first components of this sum.

sblood x; z; tð Þ ¼X

iNntissue

λiui x; zð ÞTi tð Þ

Moreover, in order to further remove any motion artifact at themicrometric scale coming from the hindlimb stimulation of the animal,we decided to apply also a numerical Butterworth high-pass temporalfilter with a cut-off frequency of 75 Hz.

Eventually, we evaluated the power Doppler signal as the modulusof the Doppler signal at each pixel.

PD x; z; tð Þ ¼ ∥sblood x; z; tð Þ∥

Computing the activation map

To obtain the activation map to detect the augmentation due to thesensory-motor activations with respect to the baseline, the PowerDoppler signal was normalized:

PDnorm x; z; tð Þ ¼PD x; z; tð Þ− 1

nt

XtPD x; z; tð ÞffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiX

tPD x; z; tð Þ− 1

nt

XtPD x; z; tð Þ

� �2r :

Then we evaluated the spatial correlation between the temporalevolutions of the power Doppler signal, PDnorm(x, z, t) with a referencefunction ref(t) representing the delivered stimulations.

ref tð Þ ¼ ∏T1

tð Þ⊗Xnδ t−nT0ð Þ

with ∏Tt

ðtÞ being a rectangular function with a width T1, δ being a Dirac

function and ⊗ stands for convolution. T0 and T1 respectively refer tothe period and the duration of the evoked tasks and n is the numberof repetition of the stimuli.

The activation map Actmap(x, z) can then be computed using theformula:

Actmap x; zð Þ ¼ maxτXt

ref norm t−τð ÞPDnorm x; z; τð Þ( )

with

ref normðtÞ ¼re f ðtÞ− 1

nt∑t re f ðtÞffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

∑t ðre f ðtÞ− 1nt∑t re f ðtÞÞ

2q :

Thresholding the activation map

The goal of this subsection is to determine if/where a reliable activa-tion is present on the activation map. To this extent we have to find athreshold over which the activation can be considered significant. Theactivated pixels are defined with the formula:

Actmap x; zð ÞNthreshold:

The threshold can be computed by studying the fluctuations of acti-vation map Actmapbaseline(x, z) computed using equation XX from base-line acquisitions where no electrical stimulus was delivered.

threshold ¼ 2 �ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiVar Actmapbaseline x; zð Þð Þ

pwith Var standing for the spatial variance.

Eventually, a spatial Pearson’s coefficient was evaluated to retrievethe likelihood of the activations of the three configurations underinvestigation.

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