A method for simulation of pulsed pressure fields from arbitrarily shaped, apodized and excited ultrasound transducers is suggested. It relies on the Tupholme-Stepanishen method for calculating pulsed pressure fields, and can also handle the continuous wave and pulse-echo case. The field is calculated by dividing the surface into small rectangles and then Summing their response. A fast calculation is obtained by using the far-field approximation. Examples of the accuracy of the approach and actual calculation times are given. >

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Calculation of pressure fields from arbitrarily shaped, apodized, and excited ultrasound transducers

Hyperthermia (HT) is used in the clinical management of cancer and benign disease. Numerous biological and clinical investigations have demonstrated that HT in the 41–45°C range can significantly enhance clinical responses to radiation therapy, and has potential for enhancing other therapies, such as chemotherapy, immunotherapy and gene therapy. Furthermore, high-temperature hyperthermia (greater than 50°C) alone is being used for selective tissue destruction as an alternative to conventional invasive surgery. The degree of thermal enhancement of these therapies is strongly dependent on the ability to localize and maintain therapeutic temperature elevations. Due to the often heterogeneous and dynamic properties of tissues, most notably blood perfusion and the presence of thermally significant blood vessels, therapeutic temperature elevations are difficult to spatially and temporally control during these forms of HT therapy. However, ultrasound technology has significant advantages that allow for a higher degree of spatial and dynamic control of the heating compared to other commonly utilized heating modalities. These advantages include a favorable range of energy penetration characteristics in soft tissue and the ability to shape the energy deposition patterns. Thus, heating systems have been developed for interstitial, intracavitary, or external approaches that utilize properties such as multiple transducer arrays, phased arrays, focused beams, mechanical and/or electrical scanning, dynamic frequency control and transducers of various shapes and sizes. This article provides a general review of a selection of ultrasound hyperthermia systems that are either in clinical use or currently under development, that utilize these advantages as a means to better localize and control HT for the aforementioned therapies.

Ultrasound technology for hyperthermia.

Computer data are presented for a nonplanar phased-array proposed as an applicator for deep, localized hyperthermia. The array provides precise control over the heating pattern in three dimensions. The array elements form a rectangular lattice on a section of a sphere so that the array has a natural focus at its geometric center when all its elements are driven in phase. When compared to a planar array with similar dimensions, the spherical-section array provides higher focal intensity gain. The relative grating-lobe level (with respect to the focus) is lower for scanned foci synthesized with this array (compared to a planar array with equal center-to-center spacing and number of elements). The spherical-section array has been simulated for use as a spot-scanning applicator as well as an applicator producing directly synthesized heating patterns. >

A spherical-section ultrasound phased array applicator for deep localized hyperthermia

Experiments were conducted to explore the potential of stabilized microbubbles for aiding tissue ablation during ultrasound therapy. Surgically exteriorized canine kidneys were irradiated in situ using single exposures of focused ultrasound. In each experiment, tip to eight separate exposures were placed in the left kidney. The right kidney was then similarly exposed, but while an ultrasound contrast agent was continually infused. Kidneys were sectioned and examined for gross observable tissue damage. Tissue damage was produced more frequently, by lower intensity and shorter duration exposures, in kidneys irradiated with the contrast agent present. Using 250-ms exposures, the minimum intensity that produced damage was lower in kidneys with microbubbles than those without (controls) in 10 of 11 (91%) animals. In a separate study using /spl sim/3200 W/cm/sup 2/ exposures, the minimum duration that produced damage was shorter after microbubbles were introduced in 11 of 12 (92%) animals. With microbubbles, gross observable tissue damage was produced with exposure intensity /spl ges//spl sim/800 W/cm/sup 2/ and exposure duration /spl ges/10 /spl mu/s. The overall intensity and duration tissue damage thresholds were reduced by /spl sim/2/spl times/ and /spl sim/100/spl times/, respectively. Results indicate that acoustic cavitation is a primary damage mechanism. Lowering in vivo tissue damage thresholds with stabilized microbubbles acting as cavitation nuclei may make acoustic cavitation a more predictable, and thus practical, mechanism for noninvasive ultrasound surgery.

Microbubble-enhanced cavitation for noninvasive ultrasound surgery

Ultrasound phased arrays offer several advantages over single focused transducer technology, enabling electronically programmable synthesis of focal size and shape, as well as position. While phased arrays have been employed for medical diagnostic and therapeutic (hyperthermia) applications, there remain fundamental problems associated with their use for surgery. These problems stem largely from the small size of each array element dictated by the wavelength employed at surgical application frequencies (2-4 MHz), the array aperture size required for the desired focal characteristics, and the number of array elements and electronic drive channels required to provide RF energy to the entire array. The present work involves the theoretical and experimental examination of novel ultrasound phased arrays consisting of array elements larger than one wavelength, minimizing the number of elements in an aperture through a combination of geometric focusing, directive beams, and sparse random placement of array elements, for tissue ablation applications. A hexagonally packed array consisting of 108 8-mm-diameter circular elements mounted on a spherical shell was modeled theoretically and a prototype array was constructed to examine the feasibility of sparse random array configurations for focal surgery. A randomly selected subset of elements of the prototype test array (64 of 108 available channels) was driven at 2.1 MHz with a 64-channel digitally controlled RF drive system. The performance of the prototype array was evaluated by comparing field data obtained from theoretical modeling to that obtained experimentally via hydrophone scanning. The results of that comparison, along with total acoustic power measurements, suggest that the use of sparse random phased arrays for focal surgery is feasible, and that the nature of array packing is an important determinant to observed performance.

Sparse random ultrasound phased array for focal surgery

A method is presented for calculation of the sound field from a rectangular continuous-wave source surrounded by a plane grid baffle. The approach is illustrated for square sources of 0.5, 1, 2, 5, 10, 20, and 100 lambda on a side. These results are compared to the sound fields produced by similarly sized circular sources. The beam widths and locations of on-axis minima are similar for the two sources, but the transverse pressure distribution is more uniform in the near-field of the square source. The effects of attenuation on the sound field of a square source are examined. >

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L.A. Frizzel

Papers

Sound field calculation for rectangular sources