Simulations and measurements of transcranial low-frequency ultrasound therapy: skull-base heating and effective area of treatment

Non-invasive ultrasonic surgery of the brain in non-human primates

High-intensity focused ultrasound causes selective tissue necrosis efficiently and safely, namely, in the prostate, liver, and uterine fibroid. Nevertheless, ablation of brain tissue using focused ultrasound remains limited due to strong aberrations induced by the skull. To achieve ultrasonic transcranial brain ablation, such aberrations have to be compensated. In this study, non-invasive therapy was performed on monkeys using adaptive correction of the therapeutic beam and 3D simulations of transcranial wave propagation based on 3D computed tomographic (CT) scan information. The aim of the study was two-fold: induce lesions in a non-human primate brain non-invasively and investigate the potential side effects. Stereotactic targeting was performed on five Macaca fascicularis individuals. Each hemisphere was treated separately with a 15-day interval and animals were sacrificed two days after the last treatment. The ultrasonic dose delivered at the focus was increased from one treatment location to the other to estimate the thermal dose for tissue alteration. Thermal doses in the brain were determined by numerical computations. Treatment efficiency and safety were evaluated histologically. The threshold for tissue damage in the brain was measured to be between 90 and 280 cumulative equivalent minutes at 43 °C. Intravenous injection of corticoids before the treatment limited the side effects.

Creating brain lesions with low-intensity focused ultrasound with microbubbles: a rat study at half a megahertz

Low-intensity focused ultrasound was applied with microbubbles (Definity, Lantheus Medical Imaging, North Billerica, MA, USA; 0.02 mL/kg) to produce brain lesions in 50 rats at 558 kHz. Burst sonications (burst length: 10 ms; pulse repetition frequency: 1 Hz; total exposure: 5 min; acoustic power: 0.47-1.3 W) generated ischemic or hemorrhagic lesions at the focal volume revealed by both magnetic resonance imaging and histology. Shorter burst time (2 ms) or shorter sonication time (1 min) reduced the probability of lesion production. Longer pulses (200 ms, 500 ms and continuous wave) caused significant near-field damage. Using microbubbles with focused ultrasound significantly reduced acoustic power levels and, therefore, avoided skull heating issues and potentially can extend the treatable volume of transcranial focused ultrasound to brain tissues close to the skull.

Transcranial passive acoustic mapping with hemispherical sparse arrays using CT-based skull-specific aberration corrections: a simulation study

The feasibility of transcranial passive acoustic mapping with hemispherical sparse arrays (30 cm diameter, 16 to 1372 elements, 2.48 mm receiver diameter) using CT-based aberration corrections was investigated via numerical simulations. A multi-layered ray acoustic transcranial ultrasound propagation model based on CT-derived skull morphology was developed. By incorporating skull-specific aberration corrections into a conventional passive beamforming algorithm (Norton and Won 2000 IEEE Trans. Geosci. Remote Sens. 38 1337-43), simulated acoustic source fields representing the emissions from acoustically-stimulated microbubbles were spatially mapped through three digitized human skulls, with the transskull reconstructions closely matching the water-path control images. Image quality was quantified based on main lobe beamwidths, peak sidelobe ratio, and image signal-to-noise ratio. The effects on the resulting image quality of the source's emission frequency and location within the skull cavity, the array sparsity and element configuration, the receiver element sensitivity, and the specific skull morphology were all investigated. The system's resolution capabilities were also estimated for various degrees of array sparsity. Passive imaging of acoustic sources through an intact skull was shown possible with sparse hemispherical imaging arrays. This technique may be useful for the monitoring and control of transcranial focused ultrasound (FUS) treatments, particularly non-thermal, cavitation-mediated applications such as FUS-induced blood-brain barrier disruption or sonothrombolysis, for which no real-time monitoring techniques currently exist.

Experimental validation of a finite-difference model for the prediction of transcranial ultrasound fields based on CT images

The prevalence of stroke worldwide and the paucity of effective therapies have triggered interest in the use of transcranial ultrasound as an adjuvant to thrombolytic therapy. Previous studies have shown that 120 kHz ultrasound enhanced thrombolysis and allowed efficient penetration through the temporal bone. The objective of our study was to develop an accurate finite-difference model of acoustic propagation through the skull based on computed tomography (CT) images. The computational approach, which neglected shear waves, was compared with a simple analytical model including shear waves. Acoustic pressure fields from a two-element annular array (120 and 60 kHz) were acquired in vitro in four human skulls. Simulations were performed using registered CT scans and a source term determined by acoustic holography. Mean errors below 14% were found between simulated pressure fields and corresponding measurements. Intracranial peak pressures were systematically underestimated and reflections from the contralateral bone were overestimated. Determination of the acoustic impedance of the bone from the CT images was the likely source of error. High correlation between predictions and measurements (R(2) = 0.93 and R(2) = 0.88 for transmitted and reflected waves amplitude, respectively) demonstrated that this model is suitable for a quantitative estimation of acoustic fields generated during 40-200 kHz ultrasound-enhanced ischemic stroke treatment.

Direct phase projection and transcranial focusing of ultrasound for brain therapy

Ultrasound can be used to noninvasively treat the human brain with hyperthermia by focusing through the skull. To obtain an accurate focus, especially at high frequencies (>500 kHz), the phase of the transmitted wave must be modified to correct the aberrations introduced by the patient's individual skull morphology. Currently, three-dimensional finite-difference time-domain simulations are used to model a point source at the target. The outward-propagating wave crosses the measured representation of the human skull and is recorded at the therapy array transducer locations. The signal is then time reversed and experimentally transmitted back to its origin. These simulations are resource intensive and add a significant delay to treatment planning. Ray propagation is computationally efficient because it neglects diffraction and only describes two propagation parameters: the wave's direction and the phase. We propose a minimal method that is based only on the phase. The phase information is projected from the external skull surface to the array locations. This replaces computationally expensive finite-difference computations with an almost instantaneous direct phase projection calculation. For the five human skull samples considered, the phase distribution outside of the skull is shown to vary by less than λ/20 as it propagates over a 5 cm distance and the validity of phase projection is established over these propagation distances. The phase aberration introduced by the skull is characterized and is shown to have a good correspondence with skull morphology. The shape of this aberration is shown to have little variation with propagation distance. The focusing quality with the proposed phase-projection algorithm is shown to be indistinguishable from the gold-standard full finite-difference simulation. In conclusion, a spherical wave that is aberrated by the skull has a phase propagation that can be accurately described as radial, even after it has been distorted. By combining finite-difference simulations with a phase-projection algorithm, the time required for treatment planning is significantly reduced. The correlation length of the phase is used to validate the algorithm and it can also be used to provide guiding parameters for clinical array transducer design in terms of transducer spacing and phase error.