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Spatio-temporal analysis of molecular delivery through the blood–brain barrier using focused ultrasound

This article has been downloaded from IOPscience. Please scroll down to see the full text article. 2007 Phys. Med. Biol. 52 5509 (http://iopscience.iop.org/0031-9155/52/18/004) View the table of contents for this issue, or go to the journal homepage for more

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PHYSICS IN MEDICINE AND BIOLOGY

Phys. Med. Biol. 52 (2007) 5509–5530

doi:10.1088/0031-9155/52/18/004

Spatio-temporal analysis of molecular delivery through the blood–brain barrier using focused ultrasound J J Choi1, M Pernot1, T R Brown1,2, S A Small3 and E E Konofagou1,2 1 2 3

Department of Biomedical Engineering, Columbia University, New York, USA Department of Radiology, Columbia University, New York, USA Department of Neurology, Columbia University, New York, USA

E-mail: [email protected]

Received 14 February 2007, in final form 11 July 2007 Published 31 August 2007 Online at stacks.iop.org/PMB/52/5509 Abstract The deposition of gadolinium through ultrasound-induced blood–brain barrier (BBB) openings in the murine hippocampus was investigated. First, wave propagation simulations through the intact mouse skull revealed minimal beam distortion while thermal deposition simulations, at the same sonication parameters used to induce BBB opening in vivo, revealed temperature increases lower than 0.5 ◦ C. The simulation results were validated experimentally in ex vivo skulls (m = 6) and in vitro tissue specimens. Then, in vivo mice (n = 9) were injected with microbubbles (OptisonTM; 25–50 µl) and sonicated (frequency: 1.525 MHz, pressure amplitudes: 0.5–1.1 MPa, burst duration: 20 ms, duty cycle: 20%, durations: 2–4 shots, 30 s per shot, 30 s interval) at the left hippocampus, through intact skin and skull. Sequential, high-resolution, T1-weighted MRI (9.4 Tesla, in-plane resolution: 75 µm, scan time: 45– 180 min) with gadolinium (OmniscanTM; 0.5 ml) injected intraperitoneally revealed a threshold of the BBB opening at 0.67 MPa and BBB closing within 28 h from opening. The contrast-enhancement area and gadolinium deposition path were monitored over time and the influence of vessel density, size and location was determined. Sonicated arteries, or their immediate surroundings, depicted greater contrast enhancement than sonicated homogeneous brain tissue regions. In conclusion, gadolinium was delivered through a transiently opened BBB and contained to a specific brain region (i.e., the hippocampus) using a single-element focused ultrasound transducer. It was also found that the amount of gadolinium deposited in the hippocampal region increased with the acoustic pressure and that the spatial distribution of the BBB opening was determined not only by the ultrasound beam, but also by the vasculature of the targeted brain region. (Some figures in this article are in colour only in the electronic version) 0031-9155/07/185509+22$30.00 © 2007 IOP Publishing Ltd Printed in the UK

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1. Introduction Current treatments of neurological and neurodegenerative diseases are limited due to the inexistence of a truly noninvasive, transient and regionally selective brain drug delivery method (Pardridge 2005). The brain is particularly difficult to deliver drugs to because of the blood–brain barrier (BBB), which is a selective and specialized vasculature that acts as a permeability barrier between the blood and brain. The impermeability of the BBB is due to tight junctions connecting adjacent endothelial cells and highly regulated transport systems of the endothelial cell membranes (Abbott et al 2006). The main function of the BBB is ion and volume regulation in order to ensure conditions necessary for proper synaptic and axonal signaling (Stewart and Tuor 1994). However, the same permeability properties that keep the brain healthy are the reason for the difficulty in the pharmacological treatment of the brain. The BBB prevents most neurologically potent drugs from entering the brain, and, as a result, has been determined as the rate-limiting factor in brain drug delivery (Pardridge 2005). Until a solution to the trans-BBB delivery problem is found, treatments of neurological diseases will remain impeded. A successful brain drug delivery method requires noninvasive, transient and regionally selective properties (Pardridge 2005). Current BBB opening methods, such as chemical modification of drugs to increase permeability and invasive neurosurgical-based transcranial drug delivery, do not satisfy these properties simultaneously (Pardridge 2005). To our knowledge, the only technique that has been shown to accommodate the necessary characteristics of a successful brain drug delivery method is focused ultrasound (FUS) with injected microbubbles. The BBB opening induced by ultrasound at or near ablation intensities has been observed while often accompanied by neuronal damage (Bakay et al 1956, Ballantine et al 1960, Patrick et al 1990, Vykhodtseva et al 1995). After reducing the acoustic intensity and duty cycle, the BBB opening was still observed, but without macroscopic damage, e.g., thermal lesions (Mesiwala et al 2002). With the addition of intravenously injected microbubbles prior to sonication, the BBB opening was determined to be transient (Hynynen et al 2001). The advantage of having microbubbles present in the circulation is that they allow for the reduction of the ultrasound intensity so that most of the disruption is contained within the vasculature with reduced likelihood of irreversible neuronal damage (Hynynen et al 2003, 2005, 2006, McDannold et al 2004, 2005, 2006, Sheikov et al 2004, 2006, Choi et al 2005, 2006, 2007). Although there are many indications that damage can be contained to minimal hemorrhage (Hynynen et al 2006), the complete safety profile of the procedure remains to be realized. In addition, indications to various mechanisms such as the dilation of vessels, temporary ischemia, mechanically induced opening of the tight junctions and the activation of various transport mechanisms have been reported (Mesiwala et al 2002, Sheikov et al 2004, 2006). However, the complexity of mechanical, thermal and physiological interactions amongst the ultrasound, microbubbles and BBB physiology make this a difficult mechanism to completely understand (Apfel 1997, Neppiras 1980). The exact mechanism of the ultrasound-induced BBB opening has not been determined yet. Fortunately, the efficacy of this method is promising with molecules up to 20 nm in diameter being successfully delivered through the BBB (Hynynen et al 2006, Choi et al 2007). The implications of successful regional and noninvasive opening of the BBB using this method cannot be overemphasized. The ultrasound-induced BBB opening may allow for the delivery of neurologically potent drugs that have been previously shelved due to low BBB permeability or lack of regional specificity. The study reported in this paper investigates the spatial and temporal variations of ultrasound-induced trans-BBB molecular delivery. All sonications were applied through

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the intact mouse skull and skin in vivo. No craniotomy was performed. Our group and others have shown that FUS can be applied transcranially (Choi et al 2005, Kinoshita et al 2006, Treat et al 2007) and with some studies characterizing the beam shape through the skull (Choi et al 2005, 2006, 2007). However, the reason for the signal intensity loss has not been elucidated yet. In this paper, using wave propagation simulations through µCT–scanned mouse skulls, it is shown that the loss in intensity is not due to the absorption from propagation through the skull, but to the reflection and refraction phenomena at the skull’s surface, i.e., the location of the highest impedance mismatch. Previous studies have shown that MRI can be used as a method to determine the BBB opening in vivo (Hynynen et al 2001). However, the studies revealed little information about the spatial and temporal variations of the model drug delivery. In this study, the molecular rate and path of diffusion were investigated in vivo using sequential, slow-contrast diffusion, high field MRI. Using the high resolution and high signal-to-noise ratio (SNR) attainable with this system, the shape of the BBB opening cannot only be determined by the shape of the ultrasound beam but also by the underlying vasculature of the sonicated region. The deposition of the model drug begins at specific sites, such as the posterior cerebral artery, before spreading to the brain parenchyma (figure 10). 2. Materials and methods In the in vivo study, the ultrasound-induced BBB opening was performed through the intact skin and skull in each mouse. However, in order to accurately characterize the transcranial ultrasonic beam, the beam distortion due to the skull was predicted using two-dimensional finite-difference model simulations. In addition, temperature rise during the BBB opening was modeled and measured in tissue specimens in vitro. Thermal deposition was computed using the bio-heat equation to determine the temperature rise using the same parameters as those used to induce the BBB opening in vivo. The microbubble effect was not incorporated in the simulations although it is known that they can enhance heat deposition (Razansky et al 2006). The purpose was to determine whether the effect of ultrasound energy deposition alone was sufficient to induce a temperature elevation large enough to cause physiological changes or thermal damage. 2.1. Simulations 2.1.1. Acoustic properties of the skull deduced from µCT images. A mouse (mass: 25 g, strain: C57BL/6, sex: male) head was severed and stabilized in a polypropylene conical tube. The tube was then placed in a µCT scanner (VivaCT 40, ScancoMedical, Inc., SWI; resolution: 10 µm thick), where three-dimensional µCT data were obtained. A single 10 µm two-dimensional coronal slice of the parietal bone was extracted from the µCT data. Although ultrasound propagation is a three-dimensional problem, the targeted skull region was found to have similar acoustic properties throughout the rest of the parietal bone region. The ultrasound beam and the targeted skull region were assumed to be axisymmetric. A threshold was applied to the µCT data and adjusted until the skull was isolated from the surrounding brain and skin tissue (figure 1(a)). The parietal bone was chosen because it was the skull region sonicated through to target the hippocampus in the in vivo study. Acoustic properties of the skull, i.e., density and speed of sound, were deduced from the skull µCT data using methods similar to those described by Aubry et al (2003) and modified for this study. The bone porosity () mapping of the skull was obtained using the linear relationship between the porosity and the Hounsfield map (Aubry et al 2003). To simplify the procedure, the lowest µCT data value of the skull was assumed to have a porosity of 1 while the highest value was set to 0

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Figure 1. Acoustic mappings of a coronal slice of the parietal bone of the mouse skull deduced from (a) a µCT image, (b) porosity (normalized), (c) density (kg m−3) and (d) speed of sound (mm µs−1) maps were used to simulate wave propagation through the skull. The bar in (a) indicates 1 mm.

(figure 1(b)). Assuming that the skull was composed of only bone and water, the mass density (ρ) was deduced from the porosity using the following linear relationship (figure 1(c)): ρ = × ρwater + (1 − ) × ρbone .


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The mass density of water (ρ water) and the maximum mass density in the cortical bone (ρ bone) were set to 1000 kg m−3 and 3170 kg m−3, respectively. The speed of sound map was extracted from the porosity using a similar linear relationship (figure 1(d)): c = cwater + (cbone − cwater ) × (1 − ). The speed of sound in water (cwater) and cortical bone (cbone) was set to 1.50 and 4.18 mm µs−1, respectively. Ultrasonic absorption has no linear relationship with porosity. However, given the minuteness of the mouse skull (∼200 µm in thickness), the details of this relationship were deemed negligible. No absorption was assumed in the wave propagation simulation. 2.1.2. Wave propagation numerical simulations. Simulations were performed with a finite-difference, wave propagation software (ACEL) developed by the Laboratoire Ondes et Acoustique, ESPCI, Paris, France (Tanter et al 1998, Aubry et al 2003). The wave propagation simulation program was based on the discretization of the linear acoustic wave equation in heterogeneous absorbing media (Aubry et al 2003): 1 ∂ 2 p(r , t) ∂ 1 ρ0 (r ) ∇ ∇p(r , t) − = S0 (r , t), 1 + τ0 (r ) ∂t ρ0 (r ) c0 (r )2 ∂t 2 where p is the acoustic pressure, τ 0 is the relaxation time, ρ 0 is the density and c0 is the speed of sound of the medium. The source term and gradient operator are S0 and ∇, respectively. The equation accurately models fluids with the speed of sound, density and absorption heterogeneities. Mode conversions and shear waves were not taken into account in the approximations, because the incidence angle of the wavefront on the skull’s interface was assumed nearly perpendicular. The details of the implementation of this equation are described by Aubry et al (2003). The acoustic mappings of the skull were loaded into ACEL and simulated ultrasound sources were placed above the mouse skull with its beam axis 2 mm away from the sagittal suture. The focal point was placed 3 mm below the top of the mouse skull so that the ultrasound beam targeted the region of the hippocampus. The ultrasound sources were placed and designed according to the characteristics of the FUS transducer used in the ex vivo and in vivo experiments, as described later in this paper. Simulated acoustic pressure data were captured in a single dimension perpendicular to the beam axis and at the focal point, with and then without taking into account the skull’s acoustic mappings. 2.1.3. Thermal deposition simulations. Pressure amplitude measurements were obtained on ex vivo skulls as described in section 2.2.2. Using these measurements, temperature changes induced by an ultrasound transducer with sonication parameters set according to the values used to induce the BBB opening in vivo (frequency: 1.525 MHz, pressure amplitude: 0.67– 1.10 MPa, burst rate: 10 Hz, burst duration: 20 ms, duration: four shots of 30 s sonication with 30 s delay between shots) were simulated. The temperature was computed using a threedimensional, finite-difference method based on the bio-heat equation (Pennes 1948, Pernot et al 2004): ∂T = ∇(k∇T ) + Q0 + Q, ρCt ∂t where ρ is the density, Ct is the specific heat capacity, T is the temperature and k is the thermal conductivity of the medium (i.e., brain tissue and water). The values were set to 1300 kg m−3, 3370 J kg−1 ◦ K−1, 37 ◦ C and 0.488 W m−1 ◦ C−1, respectively. The blood perfusion term is Q0 and the heat source term is given by α|p|2 , Q= ρc

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Figure 2. Experimental setup used in the in vivo BBB opening study.

where α is the ultrasound absorption coefficient, p is the acoustic pressure and c is the speed of sound of the medium (i.e., brain tissue and water). The ultrasound absorption coefficient was set to 7.423 Np m−1 and the speed of sound was set to 1600 m s−1. The pressure was varied according to the pressure used in our in vivo BBB opening study. 2.2. Experimental validation of simulations 2.2.1. Ultrasound equipment. Ultrasound waves were generated by a single-element spherical segment FUS transducer (center frequency: 1.525 MHz, focal depth: 90 mm, radius: 30 mm; Riverside Research Institute, New York, USA) with a circular hole (radius: 11.2 mm) in its center that held a pulse-echo diagnostic transducer (center frequency: 7.5 MHz, focal length: 60 mm), which was positioned so that the foci of the two transducers overlapped (figure 2). A cone filled with degassed and distilled water was mounted on the transducer system. The water was contained in the cone with a polyurethane membrane cap (Trojan; Church & Dwight Co., Inc., Princeton, NJ, USA). The transducer system was attached to a computer-controlled, three-dimensional positioning system (Velmex Inc., Lachine, QC, Canada). The FUS transducer was driven by a function generator (Agilent, Palo Alto, CA, USA) through a 50 dB power amplifier (ENI Inc., Rochester, NY, USA) while the diagnostic transducer was driven by a pulser-receiver system (Panametrics, Waltham, MA, USA) connected to a digitizer (Gage Applied Technologies, Inc., Lachine, QC, Canada), embedded in and controlled by a personal computer (PC). 2.2.2. Ex vivo skull experiments. Six mice (mass: 20–25 g, strain: C57BL/6, sex: male) were euthanized, and each skull was excised and degassed in saline. The bottom tip of the cone of the transducer system was partially submerged in a degassed water-filled tank. A needle hydrophone (Precision Acoustics Ltd., Dorchester, Dorset, UK; needle diameter: 0.2 mm)


(a)

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Figure 3. Location sites of FUS-targeted regions in vivo. The FUS transducer was targeted through (a) the left parietal bone of the mouse skull (red circle in (a)). The focal point was placed to overlap the left hippocampus (red circle) as seen on (b) a horizontal histology slice of the mouse brain. The right hippocampus (blue circle) was not targeted and acted as the control. After the BBB opening, a tracer was injected and monitored with (c) MRI. The horizontal slices shown in (b) and (c) were both obtained, approximately 3 mm beneath the top of the skull.

was placed at the FUS transducer focus using the three-dimensional positioning system. The dimensions of the beam were measured by performing a three-dimensional raster scan (lateral step size: 0.2 mm, axial step size: 1.0 mm) of the FUS transducer beam. The peak-to-peak pressure amplitude was measured by calculating the difference between the peak-negative and peak-positive pressure values. The lateral and axial full-width at half-maximum (FWHM) intensities of the beam were calculated to be approximately 1.32 and 13.0 mm, respectively. In a set of six measurements, each mouse skull was held stationary at 3 mm above the tip of the hydrophone. Depth calculations were performed using the pulse-echo transducer and the range equation. The skull was placed so that the FUS beam targeted through the parietal bone, approximately 2 mm away from, and perpendicular to, the sagittal and lambdoid sutures (figure 3). So, in this setup, the hydrophone was placed in the same location that the FUS transducer targeted in the in vivo study: the hippocampus. Two-dimensional pressure fields were then measured in the presence and then in the absence of each skull. 2.2.3. In vitro liver experiments. A 4 × 4 × 4 cm3 segment of fresh bovine liver was cut and degassed in saline. Liver was selected, because it has acoustic properties close to those of brain tissues (Duck 1990). The mouse brain was not used since it was smaller and less rigid, making it difficult to keep intact when handling. In addition, thermal lesions are more easily discernible in liver than in brain tissue. The liver segment was suspended in a bowl of degassed water with 2.54-cm-wide strips of tape. A rubber acoustic absorber was placed at the bottom of a glass container at a slight angle to the normal incidence plane to avoid standing waves and large reflections. Two thermocouple probes were connected to a dual input datalogger (Omega Engineering, Inc., Stamford, CT, USA), which was subsequently connected to a PC. In order to monitor the temperature rise induced by the FUS transducer, one of the thermocouples (Physitemp Instruments, Inc., Clifton, NJ, USA; type: copper-constantan, needle diameter:

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Figure 4. Experimental timeline of the in vivo BBB opening study.

0.33 mm) was placed through the lateral side of the liver, in a plane perpendicular to the FUS transducer beam axis. The temperatures of the container and liver were maintained at a constant temperature of 37 ◦ C by placing the container on a heated surface (Corning, Inc., Acton, MA, USA), while the water temperature was monitored with a second thermocouple (Omega Engineering, Inc., Stamford, CT, USA; type: copper-constantan, needle diameter: 3.18 mm). The water was routinely stirred in order to ensure uniform temperature. The FUS transducer was positioned so that its focus was placed below the surface of the liver segment and at the thermocouple tip by imaging with the diagnostic transducer in a three-dimensional raster scan. The temperature rise was induced by the FUS transducer using the same sonication parameters used in the thermal deposition simulations and in vivo experiments. 2.2.4. Comparison between transcranial simulations and ex vivo skull experiments. Relative skull attenuation values were obtained by calculating the difference between the pressure amplitude measured after propagation through the skull and the pressure amplitude measured in water only and then dividing the total by the pressure amplitude in water only. The change in FWHM intensity was calculated by calculating the difference between the FWHM intensity through the skull and that through water only and then dividing by the FWHM intensity through water only. The displacement of the focal point was found by subtracting the locations of the peaks of the foci corresponding to the two cases. The attenuation, change in the FWHM intensity and focal point displacement values of the ex vivo skull experiments were obtained by averaging over six values measured in six different skulls. 2.3. In vivo ultrasound-induced BBB opening study In vivo BBB opening experiments were performed and monitored in two steps: (1) targeting and opening of the BBB through the intact skin and skull outside the MRI scanner and (2) monitoring of the BBB opening using slow-contrast diffusion high-resolution MRI (figure 4). 2.3.1. Animal preparation. A total of nine mice (Harlan, Indianapolis, IN, USA; mass: 25 to 30 g, strain: C57BL/6, sex: male) were used. The mice were anesthetized with a mixture of ketamine (Fort Dodge Animal Health, Fort Dodge, IA, USA; concentration: 75 mg per kg of body mass) and xylazine (Ben Venue Laboratories, Bedford, OH, USA; concentration: 3.75 mg per kg of body mass). Before sonication, the fur on the top of the mouse heads was completely removed using an electric trimmer and a depilatory cream. Between sonication and MRI scanning, the mice were switched to the administration of isoflurane to simplify the long anesthesia procedure necessary for MRI scanning (Moreno


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et al 2006, Choi et al 2007). The respiration rate was monitored during all imaging procedures. A mouse was rescanned 28 h after sonication and then prepared for histology (see section 2.3.5). A 28 h delay was chosen to allow for the mouse to recuperate after the initial MRI scan. The other eight mice were euthanized immediately after MRI scanning. All procedures used on the mice were approved by the Columbia University Institutional Animal Care and Use Committee. 2.3.2. Experimental setup. The same ultrasound system and equipment setup described in the validation of simulation experiments were used for the following in vivo experiments. Each mouse was anesthetized and placed prone on a table with its head stabilized on a paper bedding (figure 2). A degassed water-filled container with the bottom made of thin acoustically and optically transparent Saran Wrap (Saran; SC Johnson, Racine, WI, USA) was placed on top of the mouse head. Ultrasound gel was used to eliminate any air trapped between the water container and the mouse skin in order to eliminate the impedance mismatch. Finally, the FUS transducer was placed in the water container with its beam axis perpendicular to the surface of the skull. The focus of the transducer was positioned in the mouse brain using a grid positioning method described in a previous study (Choi et al 2007). The beam axis of the transducer was aligned 2 mm away from the sagittal and lambdoid sutures and the focal point was placed 3 mm beneath the top of the left parietal bone of the skull. In this placement, the focus of the FUS beam would overlap with the left hippocampus and the left posterior cerebral artery (PCA) (figures 3(a) and (b)). The right hippocampus was not targeted and was used as the control. Pulsed wave FUS (burst rate: 10 Hz, burst duration: 20 ms, duty cycle: 20%) was applied through the intact skin and skull in a series of 2–4 shots. Each shot consisted of 30 s of sonication. A 30 s interval between each shot allowed the heat to dissipate. The FUS sonication procedure was performed once in each mouse brain. FUS-induced BBB opening was studied in two different sets of experiments in order to (1) roughly determine the acoustic pressure and microbubble concentration necessary to induce the BBB opening noninvasively and locally on the left hippocampus and (2) analyze the spatio-temporal nature of gadolinium deposition through BBB openings using MRI. In the first set of experiments, a total of four mice were used. Each mouse was injected intravenously (IV) via the tail vein approximately 1 min prior to sonication (figure 4) with a bolus of 50 µl of ultrasound contrast agents (OptisonTM; GE Healthcare, Chalfont St. Giles, UK) constituting of microbubbles (mean diameter: 3.0–4.5 µm, concentration: 5.0–8.0 × 108 bubbles per ml). The relatively larger amount of 50 µl than what was used before was administered because tail vein injections were difficult to perform, and it was necessary to ensure that a significant amount of microbubbles entered circulation after injection. Each mouse was then separately sonicated at different acoustic pressure settings of 0.53, 0.67, 0.80 and 1.13 MPa, in a series of four shots of 30 s in duration with 30 s delays between shots. The reported acoustic amplitudes were peak-to-peak pressure values, measured in a degassed water-filled tank and corrected using the skull attenuation values obtained in the ex vivo skull experiments. In the second set of experiments, five mice were separately injected intravenously with a bolus of 25 µl of microbubbles via the tail vein following a cut-down procedure which allowed for high confidence in the amount of OptisonTM injected. One mouse was not sonicated and acted as sham while the other mice were sonicated with 0.67 (l = 1) and 0.80 MPa (k = 3). 2.3.3. MRI acquisition. MR images were obtained with a 9.4 Tesla system (Bruker Medical; Boston, MA, USA). The mice were placed in a plastic tube with a 3.8 cm diameter birdcage

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coil attached and inserted vertically into the magnet 60–90 min after sonication. Sequential (n = 3–4) T1-weighted spin echo MR images were then obtained (repetition time/echo time (TR/TE): 246.1 ms/10 ms, bandwidth (BW): 50 505.1 Hz, matrix size: 256 × 256, field of view (FOV): 1.92 × 1.92 cm2, slice thickness: 0.6 mm, number of excitations (NEX): 5). These images were used to roughly determine whether FUS had caused any tissue damage and to check for any significant field inhomogeneities between the left and right hippocampus. Once the pre-gadolinium scans were completed, 0.5 ml of MRI contrast agent (Omniscan; Amersham Health, AS Oslo, NOR; molecular weight: 573.66), i.e., a molecule that normally does not traverse the BBB (Ford et al 1997), was administered intraperitoneally (IP) via a catheter to depict the BBB opening (Choi et al 2007). IP injection allowed for the slow uptake of the MRI contrast agent into the bloodstream (Moreno et al 2006, Choi et al 2007). After injection of the MRI contrast agent, sequential (n = 12–41) T1-weighted images were obtained for 45–180 min. The mouse sonicated at 0.67 MPa in the second set of experiments survived for 28 h post-sonication and the BBB closing was determined by imaging with sequential MRI, injecting a second dosage of gadolinium, and then imaging with sequential MRI again. 2.3.4. MRI analysis. In both sets of experiments, regions of the MRI were post-processed using four methods to analyze the contrast enhancement. In the first method, the entire image of each MRI image was subtracted by a spatially averaged 11 × 11 pixel region of the right hippocampus that was not sonicated. The left (targeted) and right (control) hippocampi were compared in each mouse and any pixel intensity value above 2.5 standard deviations of the averaged spatial region was determined to be a contrast-enhanced region, indicating a possible BBB opening. Thresholding at 2.5 standard deviations was used because it provided a statistically significant differentiation between unaffected and BBB-opened regions in all mouse experiments. In the second method, the temporal nature of the level of contrast enhancement was quantified in the second set of experiments. A 11 × 11 pixel region of the left and right striatum and a 11 × 11 pixel region of the left and right hippocampi were spatially averaged separately. The normalized increase in signal intensity for each region was measured by subtracting the left region from its equivalent right region and then dividing by the right region for each MRI image. The results from three mice tested were averaged and the standard deviations were recorded. This value was then multiplied by 100%. In this manner, the percentage increase in signal intensity of the left region over the right region was calculated. In the third method, contour maps describing the temporal nature of contrast enhancement were quantified in the second set of experiments. For each mouse, any pixel intensity value above 2.5 standard deviations of the averaged spatial striatum region was determined to be a contrast-enhanced region, indicating a possible BBB opening. This was repeated at various time points and color-coded. Finally, contour maps describing the spatial distribution of contrast enhancement were quantified in the second set of experiments. For each mouse, the regions of contrast enhancement above 2.5, 5.5, 8.5 and 11.5 standard deviations of the averaged spatial striatum regions were determined to be contrast-enhanced regions. Each level was color-coded for more easily distinguishable features. The two contour maps were finally overlaid on top of the corresponding MRI image. 2.3.5. Histology. After the last MRI sequence was obtained, the mouse sonicated at 0.67 MPa in the second set of experiments was prepared for histology in order to assess the possibility of any macroscopic damage. The mouse was anesthetized with IP-injected ketamine/xylazine and then perfused with saline and 10% neutral buffered formalin. The mouse head was


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removed and fixed in formalin. The brain was embedded in paraffin, serially sectioned into 6 µm thick horizontal slices and then stained with Haematoxylin and Eosin (H&E). 3. Results 3.1. Simulations and validation experiments The results of the one-dimensional wave propagation simulation through the skull were compared to a single dimension of the ex vivo skull experimental beam profile results (figure 5(a)). The attenuation of the pressure amplitude was 18.6% and 18.1% for the simulation and ex vivo experiments, respectively. The FWHM changes were 0.038 and 0.046 mm in the simulation and ex vivo skull experiment, respectively. There were no significant displacements of the focal point locations for both the simulation and the experiment. The results of heat deposition for both the simulations and the in vitro liver experiments were obtained and compared (figure 5(b)). Both simulation and experimental results show that the temperature rise remained under 0.5 ◦ C using the sonication parameters near the threshold to induce the BBB opening (0.67 and 0.80 MPa). 3.2. In vivo BBB opening study Contrast enhancement on the MRI images was used to monitor the opening of the BBB. Diffusion of the MRI contrast agent through BBB-opened regions resulted in an increase in pixel signal intensity on the MRI images. The FUS beam was targeted through the left parietal bone region (figure 3(a)) so that the beam overlapped with the left hippocampus (figure 3(b)). The orientation of the histology slice in figure 3(b) is similar to that of the MRI slices obtained (figure 3(c)). Figure 6 depicts MRI images from the first set of experiments with 50 µl OptisonTM intravenously injected. No significant differentiable contrast enhancement (figure 6(a)) was observed at the acoustic pressure amplitude of 0.53 MPa. However, contrast enhancement was observed at 0.67 MPa and beyond, and the area of contrast enhancement increased with pressure amplitude (figures 6(b)–(d)). It should be noted that in these sonications, the method used to target the left hippocampus was still being optimized. The accuracy was improved in the second set of experiments. At pressure amplitudes of 0.67 and 0.80 MPa, the area of contrast enhancement decreased when the amount of OptisonTM injected was reduced from 50 to 25 µl (figures 6–8). In the second set of experiments, three mice were separately intravenously injected with 25 µl OptisonTM and sonicated at 0.0 (sham), 0.67 and 0.80 MPa. The path of gadolinium deposition was tracked over 107, 98 and 180 min, respectively. There was no detectable contrast enhancement in the sham experiment. Figures 7 and 8 show how, after sonication with a pressure amplitude of 0.67 and 0.80 MPa, respectively, the gadolinium first appears in the PCA, or in its immediate vicinity, then slowly permeates throughout the surrounding regions, eventually encompassing the entire left hippocampus. Approximately 28 h after sonication at 0.67 MPa, the mouse was injected with a second equal dose of gadolinium to what was used immediately after sonication, but no contrast enhancement was seen, indicating the BBB closure (figures 8(f)–(i)). The BBB had closed sometime between 4 and 28 h after sonication. The MRI signal intensity was tracked in the striatum and hippocampus regions for over 90 min (figure 9). Contour maps of the temporal nature and spatial distribution of gadolinium deposition are depicted in figures 10 and 11. Figure 10 depicts the deposition of gadolinium along the vessels and different regions of the brain at different time points. Figure 11 depicts the spatial extent of the deposition of gadolinium. In both figures, there is clearly a non-uniform distribution of gadolinium along the vasculature path.

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Figure 5. Comparison of wave propagation and thermal deposition simulations with experiments. (a) Simulation (solid) and experimental (dotted) lateral beam profile measurements. The units were normalized to the peak intensity at the absence of the skull. The beam was well formed through the mouse skull (red) with 33% attenuation of the acoustic intensity relative to no skull (blue), with a good agreement between simulations and experiments (i.e., the lines overlap). (b) Temperature change in simulations (solid) and in vitro (dotted) experiments at three different pressure amplitudes.

4. Discussion 4.1. Simulations and validation experiments Previously developed models of transskull wave propagation and thermal deposition in the brain were modified and validated experimentally for use in mice. It has previously been shown that the FUS-induced BBB opening can be applied transcranially in mice (Choi et al


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Figure 6. In vivo dependence of contrast enhancement on pressure amplitude with 50 µl of OptisonTM injected intravenously. Four mice were each separately sonicated at pressure amplitudes of (a) 0.53 MPa, (b) 0.67 MPa, (c) 0.80 MPa and (d) 1.13 MPa. Each MR image shown was obtained approximately at 105 min post-sonication and 30 min post-gadolinium injection. Note the increase of the contrast-enhanced region with the pressure amplitude. The bar in (a) indicates 1 mm.

2005, 2006, Kinoshita et al 2006) and in rats (Treat et al 2007). Choi et al (2005, 2006, 2007) characterized the FUS beam effects at the skull and showed that the beam attenuation, change in beamwidth, focus location displacement and beam distortion did not vary significantly as a result of interaction with the mouse skull (figure 5(a)). In this paper, the effects of the skull were further investigated by showing simulation and ex vivo skull experimental findings that were in good agreement even though absorption properties of the skull were not taken into account in the simulations. This implies that absorption does not significantly contribute to attenuation, change in FWHM beamwidths, focus location displacements and beam distortions through the

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Figure 7. In vivo spatio-temporal variation of the FUS-induced BBB opening after sonication at 0.80 MPa and injection of 25 µl of OptisonTM. Sequential T1 MRI of a single mouse brain depicted the slow diffusion of intraperitoneally injected gadolinium (a) 17 min, (b) 44 min, (c) 82 min and (d) 120 min after gadolinium injection. The entire targeted (left) hippocampus is gradually infused by the gadolinium while the control (right hippocampus) remains impenetrable.

skull. Most of the intensity loss, as a result, is due to reflection and refraction phenomena at the skull surface. There was also no significant difference in these properties among different skulls. As a result of these findings, a single-element FUS transducer operating at 1.525 MHz was capable of producing a single, tight focus through the mouse skin and skull and inside the brain while the changes due to aberrations induced by the presence of the skull could be reliably predicted. According to the results of the simulations and ex vivo skull experiments


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Figure 8. In vivo detection of reversibility of the FUS-induced BBB opening after sonication at 0.67 MPa and injection of 25 µl of OptisonTM. Sequential T1-weighted MR images of the mouse brain approximately (I) 90 min after sonication depicting the BBB opening and (II) 28 h after sonication depicting the BBB closure in the left hippocampus were obtained. The right hippocampus was not FUS targeted and acted as the control in each case. All of the images are shown on the same grayscale. The BBB opening was depicted by tracking (a) before gadolinium injection and (b) 13 min, (c) 39 min and (d) 65 min after gadolinium injection. The BBB opening at 0.67 MPa was more subtle than at 0.80 MPa (figure 7). The BBB-opened region was clearly seen 159 min post-gadolinium injection. The entire image at this time point was normalized to the control hippocampus and the resulting image (e) shows clear contrast enhancement in the left hippocampus. Approximately 28 h after sonication, (f) pre-gadolinium images were obtained. The same amount of gadolinium was administered and its deposition was tracked (g) 13 min, (h) 39 min and (i) 65 min after gadolinium injection. Only in (I), where the BBB was opened, did the left hippocampus get infused with gadolinium. The ultrasound-induced BBB opening thus closed in less than 28 h.

in this study, corrections may be avoided and a well-formed beam may be assumed. This allows for the avoidance of complicated craniotomy procedures and for the reduction of the associated morbidity and mortality rates.

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Thermal deposition was studied in a theoretical model at the approximate pressure threshold (0.67 MPa) of the BBB opening in vivo. The peak temperature rise at this amplitude in both the simulation and the in vitro liver experiment was below 0.5 ◦ C (figure 5(b)), which is within the allowable temperature levels in diagnostic imaging (Barnett et al 2000). The simulations accurately predicted the temperature rise induced by the FUS transducer at 0.67 and 0.80 MPa. Deviations between the simulation and experimental data became apparent at 1.10 MPa after the first shot, most likely because heating of the tissue specimen increased the absorption of thermal energy over time. This change in absorption was not factored in our simulations. In addition, the pause between each sonication shot decreased the overall temperature elevation by allowing the heat to dissipate. Acoustic energy at the absence of microbubbles is clearly not enough to trigger a thermally induced physiological effect. However, heat deposition may be enhanced at the cellular or tissue level by microbubbles (Razansky et al 2006). Since microbubbles were not included in the simulations and in vitro liver experiments, there could still be a significant microbubble-enhanced temperature rise. This will be verified in future experiments.


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Figure 9. Temporal monitoring and quantitative analysis of the in vivo BBB opening. (a) Gadolinium deposition in the left (green) and right (cyan) hippocampi, and the left (red) and right (blue) striatum regions were monitored. The BBB in the left hippocampus was opened while the right hippocampus was not targeted and acted as the control. No striatum region was targeted. The BBB opening was monitored in the (b) striatum and the hippocampus at 0.80 MPa. No BBB opening was seen in the striatum, indicating the reliability of the MRI signal increase obtained.

4.2. In vivo BBB opening study A molecule that normally does not traverse the BBB was regionally targeted and contained to the entire left hippocampus using a single-element FUS transducer targeted through the intact skin and skull in mice. No craniotomy was required and all experiments were performed completely noninvasively. The BBB opening was determined by contrast enhancement of the MRI signal (Hynynen et al 2001). T1-weighted MRI was selected because gadolinium increased the T1 signal intensity in the region it permeated (Hynynen et al 2001, Choi et al

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Figure 10. In vivo spatio-temporal variation of the FUS-induced BBB opening after sonication at 0.80 MPa and injection of 25 µl of OptisonTM. Sequential T1 MRI of a single mouse brain depicted the slow diffusion of gadolinium 13 min (red), 25 min (blue), 38 min (yellow) and 50 min (green) after intraperitoneal injection.

Figure 11. In vivo spatial distribution of the FUS-induced BBB opening after sonication at 0.80 MPa and injection of 25 µl of OptisonTM. A single mouse brain was analyzed 50 min postgadolinium injection. Regions of contrast enhancement above 2.5 (green), 5.5 (yellow), 8.5 (blue) and 11.5 (red) standard deviations above the average MRI signal intensity of the right striatum were highlighted.


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2007). In contrast, a previous study (Choi et al 2007) showed that gadolinium repressed the T2 signal intensity and thus no new anatomical information could be extracted regarding where the gadolinium was deposited. It should be noted that a contrast-enhanced region within a voxel does not necessarily correspond to a BBB-opened region, since the MRI contrast agent could have diffused throughout the brain’s interstitial tissue during the MRI scan time. Nevertheless, the contrast enhancement describes the spatial extent of the molecular delivery and shows that the model drug was delivered to the entire hippocampus (figures 7 and 8). Figures 6–8 show that the area of localization of the contrast enhancement increased with pressure and the amount of microbubbles injected. Gadolinium was locally deposited in the left hippocampus using pressure amplitudes of 0.67 and 0.80 MPa and 25 µl OptisonTM. In these experiments, the spatial behavior of gadolinium diffusion throughout the hippocampi was monitored over an extended period of time by spatially averaging the MRI signal intensity in the hippocampus (figures 7 and 8). It was found that the MRI signal intensity of the left hippocampus increased at a higher rate at 0.80 MPa than at 0.67 MPa, indicating either that the area of the BBB opening was greater or that the effective opening of the BBB was relatively larger at 0.80 MPa. This offers the possibility of controlling the extent of drug delivery by simply adjusting the pressure amplitude applied. However, further study in this area needs to be performed before a concrete conclusion can be drawn. In addition to tracking the signal intensity level, the path of gadolinium deposition and the shape of contrast enhancement were monitored. Gadolinium began to diffuse at the PCA or within its immediate vicinity and then into the surrounding regions following a non-uniform path (figures 7, 8(a)–(e), 10 and 11). In other words, the shape of contrast enhancement when the FUS beam is targeted over a non-homogeneous vasculature does not correspond to the circular shape of the FUS beam focus (figures 10 and 11). This shape is most apparent in figure 10 at 25 min post-gadolinium injection (blue region). The irregularity of contrast enhancement was also seen in a previous study (Choi et al 2007), using pressure amplitudes greater than 2.0 MPa and a different OptisonTM concentration. There is clear indication that the larger vessels or the regions within their immediate vicinity (relative to the pixel size of less than ∼75 µm) are more sensitive to the BBB opening. This is in good agreement with what has been observed microscopically with brain arterioles exhibiting a greater BBB opening than capillaries (Sheikov et al 2006). The implication of this finding is that the brain region that the FUS beam targets will have a specific drug deposition behavior that is dependent on the vasculature in the target region. In this study, placing the FUS beam over the PCA was advantageous since it allowed for the model drug to be delivered to the entire left hippocampus. So, the question arises: what are the differences in behavior between targeting larger vessels and more homogeneous networks of capillaries? Further analysis of this behavior needs to be studied to compare the effectiveness and safety of the sonication of these different targets. Figures 8(a)–(d) depict an unexpectedly high contrast enhancement in the right control PCA. This may be because of several reasons including head movement during sonication, simple variability amongst the different mice and the condition of the mouse post-sonication. The same mouse survived, and its BBB was found to close within 28 h after sonication (figures 8(f)–(i)). The same mouse was then sacrificed, and based on high-resolution MRI and preliminary histology, no macroscopic structural damage to the hippocampus was observed (figure 12). Other than FUS, most of the currently available brain drug delivery methods in the clinic are either noninvasive or localized, but not both simultaneously (Pardridge 2005). To our knowledge, the only other technique that has the potential to noninvasively and locally open the BBB involves the chemical modification of each drug to allow the molecule to cross the

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Figure 12. H&E stained horizontal histology slice of the (a) left and (b) right hippocampi of the mouse brain depicting no macroscopic structural damage to the hippocampus. The mouse was sonicated with 0.67 MPa after IV injection of 25 µl of OptisonTM and survived for 1.5 days. This is the brain of the same mouse where MR images were depicted in figure 8. The bar indicates 500 µm.

BBB through endogenous transport systems. However, this technique would require special attention and modification of each drug and transport system resulting in significant drug development costs. In contrast, FUS-induced BBB opening has the potential to deliver drugs noninvasively, transiently and locally in the brain for a wide range of drugs. Future studies will be concerned with following the BBB opening behavior over several hours in a single MRI session to determine when and how it closes. In addition, the influence of vessel characteristics will be further studied with a detailed look into the difference in effectiveness and safety of drug delivery in different vascular regions. Finally, methods in combination with the proposed technique will be investigated to noninvasively treat neurodegenerative diseases that affect the hippocampus, such as Alzheimer’s disease. 5. Conclusion In this paper, the in vivo BBB opening phenomenon and the technical issues associated with opening the BBB in mice were studied with theoretical models as well as ex vivo and in vivo experiments. First, theoretical models of wave propagation and thermal deposition were developed and validated experimentally. It was found that the loss of acoustic energy through the mouse skull was due to reflection and refraction phenomena, not absorption. In addition, in the absence of microbubbles, the acoustic energy used to induce the BBB opening did not generate a temperature rise significant enough to trigger a physiological effect. Secondly, the spatio-temporal nature of gadolinium deposition through ultrasound-induced BBB openings was studied in vivo in mice using high-resolution MRI, which allowed for detailed spatial and temporal analysis of the drug behavior for up to 28 h. Using this technology, it was found that the type of vasculature in the region targeted influences how the drug will be delivered, with the larger vessels or their immediate vicinity being more sensitive to the BBB opening.


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Acknowledgments This study was supported by a Special Development Award from the Whitaker Foundation. The authors especially wish to thank the Riverside Research Institute for providing the transducers that were used in this study; Mickael Tanter, PhD (Laboratoire Ondes et Acoustique, Paris, France) for providing the ACEL wave propagation software; Barclay Morrison III, PhD (Department of Biomedical Engineering, Columbia University) for help with histology and fruitful discussions; Kenneth Hess and Fan Hua (Department of Neurology, Columbia University Medical Center) for animal assistance and training; Shougang Wang, PhD (Department of Biomedical Engineering, Columbia University), Andrew Hsu, Marina Daskalopoulos and Eugenia Kwon for assistance with the in vivo mouse experiments.

References Abbott N J, Ronnback L and Hansson E 2006 Astrocyte-endothelial interactions at the blood-brain barrier Nat. Rev. Neurosci. 7 41–53 Apfel R E 1997 Sonic effervescence: a tutorial on acoustic cavitation J. Acoust. Soc. Am. 101 1227–37 Aubry J F, Tanter M, Pernot M, Thomas J L and Fink M 2003 Experimental demonstration of noninvasive transskull adaptive focusing based on prior computed tomography scans J. Acoust. Soc. Am. 113 84–93 Bakay L, Ballantine H T Jr, Hueter T F and Sosa D 1956 Ultrasonically produced changes in the blood-brain barrier AMA Arch. Neurol. Psychiat. 76 457–67 Ballantine H T Jr, Bell E and Manlapaz J 1960 Progress and problems in the neurological applications of focused ultrasound J. Neurosurg. 17 858–76 Barnett S B, Ter Haar G R, Ziskin M C, Rott H D, Duck F A and Maeda K 2000 International recommendations and guidelines for the safe use of diagnostic ultrasound in medicine Ultrasound Med. Biol. 26 355–66 Choi J J, Pernot M, Small S A and Konofagou E E 2005 Feasibility of transcranial, localized drug-delivery in the brain of Alzheimer’s-model mice using focused ultrasound Ultrasonics Symp. IEEE (Rotterdam, The Netherlands) pp 988–91 Choi J J, Pernot M, Small S A and Konofagou E E 2007 Noninvasive, transcranial and localized opening of the blood-brain barrier using focused ultrasound in mice Ultrasound Med. Biol. 33 95–104 Choi J J, Small S A and Konofagou E E 2006 Optimization of ultrasound-induced transient blood-brain barrier opening in mice Ultrasonics Symp. IEEE (Vancouver, Canada) Choi J J, Wang S, Morrison B and Konofagou E E 2007 Molecular delivery and microbubble dependence study of the FUS-induced blood-brain barrier opening in vivo Ultrasonics Symp. IEEE (New York, NY, USA) Duck F A 1990 Physical Properties of Tissue: A Comprehensive Reference Book (London: Academic) Ford J C, Wood A K, Van Winkle T J and Kundel H L 1997 Magnetic resonance imaging observations of bloodbrain-barrier permeability in an animal model of brain injury Acad. Radiol. 4 115–26 Hynynen K, McDannold N, Sheikov N A, Jolesz F A and Vykhodtseva N 2005 Local and reversible bloodbrain barrier disruption by noninvasive focused ultrasound at frequencies suitable for trans-skull sonications Neuroimage 24 12–20 Hynynen K, McDannold N, Vykhodtseva N and Jolesz F A 2001 Noninvasive MR imaging-guided focal opening of the blood-brain barrier in rabbits Radiology 220 640–6 Hynynen K, McDannold N, Vykhodtseva N and Jolesz F A 2003 Non-invasive opening of BBB by focused ultrasound Acta Neurochir. Suppl. 86 555–8 Hynynen K, McDannold N, Vykhodtseva N, Raymond S, Weissleder R, Jolesz F A and Sheikov N 2006 Focal disruption of the blood-brain barrier due to 260-kHz ultrasound bursts: a method for molecular imaging and targeted drug delivery J. Neurosurg. 105 445–54 Kinoshita M, McDannold N, Jolesz F A and Hynynen K 2006 Targeted delivery of antibodies through the blood-brain barrier by MRI-guided focused ultrasound Biochem. Biophys. Res. Commun. 340 1085–90 McDannold N, Vykhodtseva N and Hynynen K 2006 Targeted disruption of the blood-brain barrier with focused ultrasound: association with cavitation activity Phys. Med. Biol. 51 793–807 McDannold N, Vykhodtseva N, Jolesz F A and Hynynen K 2004 MRI investigation of the threshold for thermally induced blood-brain barrier disruption and brain tissue damage in the rabbit brain Magn. Reson. Med. 51 913–23 McDannold N, Vykhodtseva N, Raymond S, Jolesz F A and Hynynen K 2005 MRI-guided targeted blood-brain barrier disruption with focused ultrasound: histological findings in rabbits Ultrasound Med. Biol. 31 1527–37

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Mesiwala A H, Farrell L, Wenzel H J, Silbergeld D L, Crum L A, Winn H R and Mourad P D 2002 High-intensity focused ultrasound selectively disrupts the blood-brain barrier in vivo Ultrasound Med. Biol. 28 389–400 Moreno H, Hua F, Brown T and Small S 2006 Longitudinal mapping of mouse cerebral blood volume with MRI NMR Biomed. 19 535–43 Neppiras E A 1980 Acoustic cavitation Phys. Rep. 61 159–251 Pardridge W M 2005 The blood-brain barrier: bottleneck in brain drug development NeuroRx 2 3–14 Patrick J T, Nolting M N, Goss S A, Dines K A, Clendenon J L, Rea M A and Heimburger R F 1990 Ultrasound and the blood-brain barrier Adv. Exp. Med. Biol. 267 369–81 Pennes H H 1948 Analysis of tissue and arterial blood temperatures in the resting human forearm J. Appl. Physiol. 1 93–122 Pernot M, Aubry J F, Tanter M, Andre F and Fink M 2004 Prediction of the skull overheating during high intensity focused ultrasound transcranial brain therapy Ultrasonics Symp. IEEE (Montreal, Canada) Razansky D, Einziger P D and Adam D R 2006 Enhanced heat deposition using ultrasound contrast agent—modeling and experimental observations IEEE Trans. Ultrason. Ferroelectr. Freq. Control 53 137–47 Sheikov N, McDannold N, Jolesz F, Zhang Y Z, Tam K and Hynynen K 2006 Brain arterioles show more active vesicular transport of blood-borne tracer molecules than capillaries and venules after focused ultrasound-evoked opening of the blood-brain barrier Ultrasound Med. Biol. 32 1399–409 Sheikov N, McDannold N, Vykhodtseva N, Jolesz F and Hynynen K 2004 Cellular mechanisms of the blood-brain barrier opening induced by ultrasound in presence of microbubbles Ultrasound Med. Biol. 30 979–89 Stewart P A and Tuor U I 1994 Blood-eye barriers in the rat: correlation of ultrastructure with function J. Comp. Neurol. 340 566–76 Tanter M, Thomas J L and Fink M 1998 Focusing and steering through absorbing and aberrating layers: application to ultrasonic propagation through the skull J. Acoust. Soc. Am. 103 2403–10 Treat L H, McDannold N, Vykhodtseva N, Zhang Y, Tam K and Hynynen K 2007 Targeted delivery of doxorubicin to the rat brain at therapeutic levels using MRI-guided focused ultrasound Int. J. Cancer 121 901–7 Vykhodtseva N I, Hynynen K and Damianou C 1995 Histologic effects of high intensity pulsed ultrasound exposure with subharmonic emission in rabbit brain in vivo Ultrasound Med. Biol. 21 969–79