WO1998007373A9 - Methods and apparatus for delivery of noninvasive ultrasound brain therapy through intact skull - Google Patents

Abstract

Methods and apparatus for delivery of ultrasound to the brain, without requiring removal of portions of the skull, call for transmission of ultrasound with a plurality of transducers (12) aimed to induce cavitation at least at a selected region of the brain. An excitation source (22) is arranged for driving at least selected transducers at differing phases with respect to one another, e.g., to compensate for phase shifts (or phase distortions) effected by the skull on the ultrasound output by each transducer. As a result, the ultrasound waves reaching the selected region from the transducers arrive substantially in phase with one another, e.g., within 90 degrees and preferably 45 degrees, and still more preferably 20 degrees of one another.

Description

Methods and Apparatus for Delivery of Noninvasive Ultrasound Brain Therapy Through Intact Skull

Sponsorship The research resulted, at least, in part, from work performed under NCI

Research Grant No 46627.

Reference to Related Applications

This application claims the benefit of the filing date of and is a continuation- in-part of copending, commonly assigned United States Patent Application Serial No. 08/711,289, filed August 21, 1996 (Attorney Docket No. 0092664-0008), the teachings of which are incorporated herein by reference. This application also claims the benefit of the filing date of copending United States Provisional Application Serial Nos. 60/034,084 (filed 12/23/96), and 60/045,453 (filed 5/1/97). The teachings of those provisional applications are incorporated herein by reference.

Background of Invention

The invention pertains to medical systems and, more particularly, to methods and apparatus for non-invasive application of focused ultrasound to the brain. The invention can be used, for example, in the diagnosis and treatment of neural ailments.

According to the prior art, treatment of tissues lying at specific locations within the skull are limited to removal or ablation. While these treatments have proven effective for certain localized disorders, such as tumors, they involve delicate, time-consuming procedures that may result in destruction of otherwise healthy tissues. The treatments are generally not appropriate for disorders in which diseased tissue is integrated into healthy tissue, except in instances where destruction of the latter will not unduly effect neurologic function.

The noninvasive nature of ultrasound surgery has special appeal in the brain where it is often desirable to destroy or treat deep tissue volumes without disturbing the healthy tissues. Focussed ultrasound beams have been used for noninvasive surgery in many other parts of the body. Ultrasound penetrates well through soft tissues and, due to the short wavelengths (1.5 mm at 1 MHz), it can be focused to spots with dimensions of a few millimeters. By heating tumorous or cancerous tissue in the abdomen, for example, it is possible to ablate the diseased portions without significant damage to surrounding healthy tissue.

Notwithstanding the potential benefits of ultrasound diagnostics and therapy of the brain, it has been commonly accepted that ultrasound cannot be applied through the intact skull. Early experiments, for example, showed that ultrasound is strongly attenuated by bone, and that brain tissue damage close to the skull results from the high temperatures caused by the energy loss. Accordingly, all of the ultrasound brain treatments performed so far have required the skull bone to be removed prior to the sonication. This makes the procedure invasive, and expensive with an added risk of complications.

The requirement of surgical removal of the skull has been the main obstacle that has prevented ultrasound therapy to be widely tested in the brain despite the possibility of its clear benefits compared to other techniques. Accordingly, an object of this invention is to provide improved medical methods and apparatus, for diagnosis and therapy of the brain. A more particular object of the invention is to provide improved methods and apparatus for application of ultrasound to the brain.

A more particular object of the invention is to provide such methods and apparatus as do not require removal of portions of the skull, via craniectomy or other such procedures.

Still another object of the invention is to provide such methods and apparatus as can be used to precisely target regions within the brain.

Still yet another object of the invention is to provide such methods and apparatus as can be used to effect heating or other physiologic change at such precisely targeted regions, without effecting substantial change in the surrounding, or other, regions of the brain or skull. Another object of the invention is to provide such methods and apparatus as can be utilized over a wide range of ultrasonic frequencies.

Still another object of the invention is to provide such methods and apparatus as can be implemented utilizing conventional materials.

Yet still another object of the invention is to provide such methods as can be implemented without excessive expense.

Summary of the Invention

The foregoing and other objects are met by the invention, which provides in one aspect methods and apparatus for delivery of cavitating ultrasound to the brain, without requiring removal of portions of the skull.

Thus, in one aspect, the invention provides an apparatus for delivering ultrasound, through intact skull, to the brain comprising a plurality of transducers and an excitation source for driving each to induce cavitation at least at a selected region of the brain. The excitation source is particularly arranged for driving at least selected transducers at differing phases with respect to one another, e.g., to compensate for phase shifts (or phase distortions) effected by the skull on the ultrasound output by each transducer. As a result, the ultrasound waves reaching the selected region from the transducers arrive substantially in phase with one another, e.g., within 90° and, preferably, within 45° and, still more preferably, within 20° of one another.

The excitation source drives the transducers to deliver ultrasound to the selected region at a frequency ranging from 0.01 MHz to 10 MHz and, preferably, from 0.1 MHz to 2 MHz. Sonication duration for the ultrasound ranges, according to further aspects of the invention, from 100 nanoseconds to 30 minutes. According to still further aspects, the invention provides for delivery of ultrasound to the selected region with continuous wave operation or burst mode operation, where burst mode repetition varies from 0.01 Hz to 1 MHz. Further aspects of the invention provide an apparatus as described above, in which only a single transducer is used.

Still further aspects of the invention provide methods for operating transducer arrays as described above.

These and other aspects of the invention are evident in the drawings and in the text that follows.

Brief Description of the Drawings

A further understanding of the invention may be attained by reference to the drawings, in which:

Figure 1 depicts an embodiment of the invention and an experimental setup for testing it.

Figure 2 depicts an embodiment of the invention for application of ultrasound to the brain of an animal.

Figure 3 depicts a phased array for application of ultrasound to the brain in accord with one practice of the invention.

Figures 4A-4H illustrate the ultrasound pressure amplitude distribution in water across the focus of a transducer according to the invention at various frequencies, with and without skull sections in front of the transducer.

Figures 5 A and 5B illustrate the effect of applying ultrasound in accordance with the invention to brain tissue.

Figures 6 A and 6B illustrate phase errors measured at the focus of ultrasound transducer arrays with a piece of skull in front of the transducers. Figures 7A-7C illustrate the pressure amplitude profiles across the focus of an ultrasound transducer phased array in water, through the bone, and through the bone with phase correction.

Figure 8 illustrates the pressure amplitude distribution along the central axis of an ultrasound transducer array without and with the phase correction.

Figures 9A-9C illustrate the ultrasound pressure amplitude distribution measured across the focus of an ultrasound phased array in water, through skull without phase correction, and through skull with phase correction.

Figure 10 depicts an embodiment of the invention for delivery of cavitating ultrasound to a patient's brain through the skull using a multi-element transducer array.

Figure 11 depicts a method for delivery of cavitating ultrasound to a patient's brain through the skull using a transducer array.

Detailed Description of the Illustrated Embodiment Discussed below are methods and apparatus according to the invention for noninvasive delivery of ultrasound through intact skull to the brain. These permit ultrasound propagation through skull to effect cavitation, without causing undesired heating of the brain or, more generally, central nervous system (CNS) tissues. These also deliver adequate ultrasound power to ablate tissues, or to otherwise induce changes, at focal points (or regions) within the brain. As used herein, "tissue" refers to fluids, tissues or other components on or within a patient's body.

Figure 10 depicts an apparatus according to the invention for delivery of ultrasound to the brain. The apparatus 10 includes an array of transducers 12 disposed on or near the external surface of the head of a human patient. The array 12 can constitute a single transducer, e.g., a spherically curved piezoelectric bowl of the type described below, though preferably, array 12 comprises a plurality of transducers arranged in a one-, two- or three-dimensional configuration.

Referring to Figure 3, for example, in one embodiment of the invention, array 12 comprises 60 individual piezoelectric ceramic transducers mounted in a bowl of circular cross-section. The transducer elements, which can be, for example, 1 cm² piezoelectric ceramic pieces, are mounted in silicone rubber or any other material suitable damping agent for minimizing the mechanical coupling therebetween. Transducer arrays of this type are known in the art, as described, for example in Fan et al, "Control of the Necrosed Tissue Volume During Noninvasive Ultrasound

Surgery Using a 16-Element Phased Array," Medical Physics, v. 22, pp. 297 et seq (1995), the teachings of which (e.g., at Figure 1 and the accompanying text) are incorporated herein by reference. The construction of a spherically curved phased array comprising multiple square-element transducers is shown in Figure 1 and the accompanying text of that publication.

In the illustrated embodiment, each transducer of array 12 is independently driven by power and the control elements 18-22 to generate ultrasound for transmission through the patient's skull into the CNS tissues. More particularly, the transducers in array 12 are individually coupled, via coaxial cables 16, to separate channels of a driving system 18. Each channel of that system 18 includes an amplifier and a phase shifter, as shown. A common radio frequency (RF) signal is driven to each channel by radio frequency generator 22. Together, the radio frequency generator 22 and driving system 18 drive the individual transducers of array 12 at the same frequency, but at different phases, so as to transmit a focused ultrasound beam through the patient's skull to a selected region within the brain. Unlike prior art systems, there is no need to remove portions of the skull beneath the array 12, e.g., via craniectomy or other such surgical procedure.

The radio frequency generator 22 can be of any commercially available type.

A preferred such generator is available from Stanford Research Systems, Model DS345. The generator is operated in a conventional way so as to generate an excitation signal, which is amplified and phase-shifted by the individual channels of driving system 18, in order to induce the corresponding transducers of array 12 to radiate ultrasound (e.g., in the range 0.01 MHz to 10 MHz).

As illustrated, each channel in the driving system 18 includes a radio frequency amplifier. These can be any RF amplifiers of the type commercially available in the art.

The phase shifting component of each channel of driving system 18 is constructed and operated in the conventional manner known in the art. Particularly, each phase shifter shifts the phase of an incoming RF excitation signal, received from RF generator 22, by an amount α,, ₂, α₃, etc., as shown in the drawing. These phase shift factors α,, α₂, ₃, etc., can be pre-stored in the channels of driving system 18 or, preferably, generated by a controller 20. That controller 20 can be a general purpose, or special purpose, digital data processor programmed in a conventional manner in order to generate and apply phase shift factors in accord with the teachings hereof.

The phase shift factors, α α₂, α₃, etc. serve two purposes. The first is to steer the composite ultrasound beam generated by transducer array 12 so that it is focused on a desired region within the patient's brain. The component of each phase shift factor associated with steering is computed in the manner known in the art for steering phased arrays. See, for example, Buchanan et al, "Intracavitary Ultrasound Phased Array System," IEEE Transactions Biomedical Engineering, v. 41, pp. 1178- 1187, the teachings of which are incorporated herein by reference. Array steering, or focusing, is particularly discussed in that article, for example, at pages 1179-1181 and, more particularly, in the section entitled "Focusing Techniques," the teachings of which are incorporated herein by reference.

The second component of each phase shift factor α,, α₂, α₃, etc., compensates for phase distortion effected by the skull in the ultrasound ouput by each transducer. In other words, the second component of the phase shift factors compensates for perturbations and distortions introduced by the skull, the skin/skull interface, the dura matter/skull interface, and by variations in the skull thickness. As those skilled in the art will appreciate, the two components that make up the phase shift factor for each channel of the driving system 18 are summed in order to determine the composite phase shift factor for the respective channel.

The phase corrections that constitute the aforementioned second component of each phase shift factor can be determined a number of ways. In one embodiment of the invention, that component is determined from measurements of the thickness of the patient's skull under each transducer in array 12. Such skull thickness measurements can be made using conventional imaging techniques, such as computed tomography (CT) or magnetic resonance imaging (MRI).

In an alternative embodiment, the aforementioned second component of each phase shift factor is determined by placing the array 12 on the patient's head and exciting individual transducers with a short ultrasound pulse. The echo back from the inner surfaces of the skull are monitored by the transducer array 12. The effect of the skull on ultrasound generated by each transducer is determined from those echos in accord with conventionally known relations.

In still further alternative embodiments, the aforementioned second component of each phase shift factor is determined by implanting small hydrophones in the patient's brain. These are used to monitor the phase of the ultrasound generated by each transducer, e.g., in a manner similar to that described below in connection with Figure 1.

In lieu of illustrated components 18-22, the transducer array 12 can be driven by a driving system of the type disclosed in Buchanan et al, supra e.g. at Figure 2 thereof, the teachings of which are incorporated herein by reference. Such a driving system would, of course, require modification in accord with the teachings hereof in order to incorporate phase shift factors α,, α₂, α₃, etc., having first and second components as described herein and above.

Referring to Figure 11, the system 10 is operated as described below in order to deliver ultrasound through the patient's skull to induce cavitation at a desired region of the brain.

In step 24, the transducer array 24 is positioned on the patient's head. This is preferably accomplished in the conventional manner known in the art for insuring ultrasound transmission to the brain. The array is typically positioned over, and as close to, the region in which cavitation is to be induced. However, where intervening or adjacent cranial or CNS tissues might be adversely affected, the array can be positioned elsewhere and focused accordingly.

In step 26, the aforementioned second component of the phase shift factor for each transducer is determined. This is accomplished in the manner described above, e.g., by individual exciting each element of the array and measuring the echo back. The alternative mechanisms described above can also be used to determine those components. Those skilled in the art will appreciate that in instances where the alternative mechanisms are used, they need not be performed after the array is positioned but, can be performed at some other prior time.

In step 28, the remaining components of each transducers' phase shift factor are determined. Particularly, those components associated with steering the array for delivery of ultrasound to the desired region are determined. Such determination is made, as indicated above, in the conventional manner known in the art for steering phased arrays.

In step 30, the array is excited, e.g., by control and driving elements 18-22, to focus ultrasound in the patient's head. As noted throughout, because the invention provides correction for phased distortion induced by the skull, that ultrasound can be supplied directly through the skull without the need for removal of a piece thereof. The ultrasound is applied in doses and timing sufficient to induce cavitation in the desired region, which may be, e.g., from 1 mm³ - 1 cm³, or larger. Those skilled in the art will appreciate that ultrasound waves in the frequency range of 0.01 MHz to 10 MHz and, preferably, from 0.10 MHz to 2 MHz can be applied with sonication duration ranging from 100 nanoseconds to 30 minutes, with continuous wave or burst mode operation. The burst mode repetition varies from 0.01 Hz to 1 MHz.

In embodiments where the transducer array 12 includes only a single transducer, e.g., a 10 cm diameter piezoelectric ceramic element as described elsewhere herein, step 26 is not utilized. In such an embodiment, the "array" is aimed based on its focal point. This is determined as a function of the size, radius of curvature and frequency output of the transducer in the manner known in the art. In a preferred embodiment, these factors are adjusted so that the transducer can be placed directly on the patient's skull, as above. However, where minor corrections are necessary, the transducer can be spaced apart from the skull, as necessary, in order to insure proper positioning of the focal point.

Theory

When an ultrasound beam propagates to a deep target location in the brain part of the energy is reflected back at the skin-skull interface due to the high acoustic mismatch between these two tissues. The propagating wave in the skull suffers attenuation losses due to scattering and absorption. The acoustic mismatch at the bone-dura interface causes part of the remaining wave to reflect back to the skull.

The total insertion loss through skull depends on the frequency, and can be, on average, about 10 dB, and 20 dB at 0.5, and 1.5 MHz, respectively. The wave is further attenuated by absorption (about 5 Np/m/MHz) while it travels through the brain to the target volume.

An ultrasound beam delivered to the brain can effect change in CNS tissues and fluids (herein, simply "tissue" or "brain tissue," etc.) by two mechanisms: heating and cavitation. The ultrasound beam can heat the tissue temperature due to energy absorption from the wave resulting in different degrees of thermal damage to the tissue depending on the temperature reached. For exposures of a few seconds, temperatures of above about 60° C are adequate to coagulate proteins and thus, necrose the tissue. The induced temperature elevation during short ultrasound exposures depends mainly on the absorbed power (<q>) although the shape and size of the focal spot can have a significant impact due to thermal conduction. The rate of temperature rise (dT/dt) at the very beginning of an ultrasound pulse can be calculated from the pressure amplitude of the field (P), as follows:

dT/dt = <q>/ p c, [1]

where

<q> = α P²/ pv [2] p is the density (p_b = 1030 kg/m³ , p_s = 1380-1810 kg/m^J) (s= skull, b = brain), c is the specific heat of the medium (c_b =3.9 kJ/kg/°C ; c_s =2.1-2.7 kJ7kg/°C), v is the speed of sound, and is the amplitude absorption coefficient of the tissue (α_b= 5Np/m/MHz, α_s = 50 Np/m at 0.5 MHz, α_s = 300 Np/m at 1.5 MHz (if all attenuated energy is assumed to be absorbed)).

In order to achieve the same temperature within the target volume as in the skull, the ultrasound beam has to be focused to overcome the difference in the acoustic properties. The square of the pressure amplitude (P²) is directly proportional to the ultrasound beam area allowing the required area gain (AG) to be calculated from Equations [1] and [2] by making the rates of temperature rise equal in the skull and the brain:

AG = (P_b7P_s ²) = (a a_h)(pjp,γ(vjvχ ,) = 30 at 0.5 MHz [3]

60 at 1.5 MHz In addition the area gain has to compensate for the energy loss due to the skull and attenuation in the brain between the skull and the focal point:

AG'= (α, e-^2pfit)-' = 13 at 0.5 MHz [4] = 250 at 1.5 MHz

where, α, is the insertion loss of skull, β is the amplitude attenuation coefficient, f is the frequency and x is the depth in the brain.

The total area gain is the product of these two area gains and is approximately 400 and 15000 at 0.5 MHz and 1.5 MHz, respectively when the focus is located at the depth of 6 cm in the brain. These calculations are first order estimates and do not take into account phase shifts introduced by the variable thickness of the bone or thermal conduction and perfusion effects.

The second mechanism, cavitation requires negative pressure amplitudes that are large enough to form gas bubbles in the tissue. The pressure wave causes the bubbles to expand and then collapse. The collapse of the bubbles causes high temperatures and pressures that can cause direct mechanical damage to the tissue. Cavitation can offer more therapeutic options than thermal exposures of brain. The cavitation threshold in the soft tissues and in bone appears to be similar.

Thus, focusing a cavitation-inducing ultrasound beam overcomes the attenuation losses in the bone and brain, but need not overcome differences in absorption coefficients, as is the case with the heat-inducing exposures. The beam area of cavitation-inducing ultrasound propagating through the skull has to be about 13 and 250 times larger than the focal area at frequencies of 0.5 and 1.5 MHz, respectively. These area gains are 30 and 60 times smaller than the gains required for induction of thermal effects. The cavitation is not influenced by thermal conduction or perfusion effects. Therefore, it is clear that cavitation has significant advantages over the thermal effects. This is particularly true in instances where the ultrasound energy must be delivered to small focal regions that require high frequencies.

Cavitation requires high pressure amplitudes but only short exposure durations, therefore cavitational effects can be induced without significant temperature elevation. For CNS tissues, sonications with durations of only 1 ms are adequate for bubble formation. The required peak intensities at 0.936 MHz during these sonications are measured to be around 4000 Wcnr² and 2000 Wcnr²at 1 ms and 1 s exposures, respectively. Using these intensity values and the average ultrasound attenuation in brain of 5 Np/m and in bone 120 Np/m at 1 MHz, the maximum peak temperature elevation in the brain can be estimated (from equations 1 and 2) to be about 60° C and 0.1° C during the 1 s and 1 ms exposures, respectively. The corresponding temperature elevations in bone, if the focus was in the bone, are 1800° C and 3.6° C. During the sonications the temperature elevation in the bone would be reduced proportionally with the area gain. These values are frequency dependent. For example, bone heating would be about 13° C for 1 ms pulse at 1.5 MHz. This short thermal exposure is below the threshold for tissue damage. Thermal exposures can be further reduced using multiple pulses that can be repeated at a low frequency (for example 0.1 Hz) thus, eliminating a temperature build up.

Examples

In one exemplary embodiment of the invention, which is illustrated in Figure 1, ultrasound was generated using a one-element transducer array having a spherically curved 10 cm diameter piezoelectric ceramic (PZT4) bowl mounted in a plastic holder using silicon rubber. The ceramic had silver or gold electrodes both on the front and back surface. The electrodes were attached to a coaxial cable that was connected to a LC matching network that matched the electrical impedance of the transducer and the cable to the RF amplifier output impedance of 50 ohm and zero phase. The matching circuit was connected to an RF- amplifier (both ENI A240L and A500 were used in the tests). The RF signal was generated by a signal generator (Stanford Research Systems, Model DS345).

The ultrasound pressure wave distributions were measured using needle hydrophones (spot diameter 0.5 and 1 mm) and an amplifier (Precision Acoustics Ltd). The amplified signal was measured and stored by a oscilloscope (Tektronix, model 243 IL ). The hydrophone was moved by stepper motors in three dimensions under computer control. The pressure amplitudes measured by the oscilloscope were stored by the computer for each location.

A piece of human skull (top part of the head: front to back 18 cm and maximum width 12 cm) was obtained and fixed in formaldehyde. The acoustic properties of formaldehyde fixed skull and a fresh skull are almost identical. The ultrasound transducer under test was positioned in a water tank the walls and bottom of which were covered by rubber mats to reduce ultrasound reflections. The tank was filled with degassed deionized water. The hydrophone was connected to the scanning frame, and positioned at the focus of the ultrasound field.

Utilizing the setup shown in Figure 1, the embodiment was tested at four different ultrasound frequencies: 0.246 MHz, 0.559 MHz, 1 MHz, and 1.68 MHz. The maximum peak pressure amplitudes achievable through the skull at the focus of the transducer was measured at each frequency. A shock wave hydrophone (Sonic Technologies Inc, ) was positioned at the acoustic focus. Bursts of 10-20 cycles were used to separate the acoustic signal from the electrical interference that was picked up by the hydrophone during sonication. Results of the testing are shown in Figures 4A- 4H.

Particularly, Figure 4A illustrates the ultrasound pressure amplitude distribution in water at the focal point of the single transducer driven at 0.246 MHz, without the skull section in place. Figure 4B illustrates this same distribution when the skull was positioned in front of the transducer as illustrated in Figure 1. Figures 4C and 4D illustrate the same distributions (i.e., with and without the skull section in place) for a frequency of 0.559 MHz. Figures 4E and 4F illustrate the same distributions for a frequency of 1 MHz. Figures 4G and 4H illustrate the same distributions for a frequency of 1.68 MHz.

To demonstrate that the single-element array of this embodiment delivers sufficient energy to induce tissue damage through the skull, in vivo rabbit experiments were performed. In these experiments a window of about 15x15 mm was created in the top of the skull. The skin was placed over the skull opening and the animal was allowed to recover. A minimum of two weeks after the surgery the animal was anesthetized again and placed on top of a sonication tank as illustrated in figure 2 and the 0.556 MHz transducer was aimed such than the focus was located at 10-15 mm in the brain. The skull piece was positioned in between the traducer and the animal. A thermocouple probe (0.05 mm constantan and copper wires were soldered together at the tip) was placed on the skull bone (on the side of the transducer that is expected to be the hottest location) under a thin layer of connective tissue that was still attached on the skull. Then 10 sonications at the maximum power level for the duration of 0.2 s were repeated with the rate of 1 Hz. The animal position was moved and the sonication repeated four times in the same location with a delay of about 5 min between the sonications to allow the bone temperature to return to the baseline. During the 10 s of pulsed sonication the bone temperature increased from the baseline of about 30° C to maximum of 43° C with rapid decay. After the sonications the rabbit was taken to a MRI scanner and Tl, T2 and contrast enhanced scan were performed. After the imaging the animal sacrificed.

Figure 5 A is a scan of the rabbit brain illustrating the effect of 10 sonications for the duration of 0.2 seconds, with a pressure amplitude of 8 MPa, repeated at a rate of 1. Hz. During the 10 seconds of pulse sonication, the bone temperature in the rabbit skull peaked from a baseline of about 30° C to a maximum of 43° C, with rapid decay. The Figure is a T2-weighted fast spin echo image across the brain. The arrow in the Figure shows tissue damage at the focal point of the transducer. The skull window on the top of the head is facing down and, thus, the ultrasound beam propagated from bottom up. Figure 4B is identical to Figure 4A, except insofar as it shows the results where the above sonication was repeated four times.

Further embodiments of the invention utilize multi-element phase arrays of the types illustrated in Figures 3 and 10, in lieu of a single transducer. By controlling the phase of the ultrasound wave as a function of transducer location, these embodiments eliminate the phase distortion caused by the skull and thus, allow accurate aiming and use of higher frequencies, thus, permitting application of ultrasound to induce cavitation through the intact skull in regions of 1 mm to 1 cm³.

Two phased arrays comprising these further embodiments had similar structure and the same driving hardware; the resonant frequency being their only significant difference. The two arrays operated at 0.6 MHz and 1.58 MHz. The radius of curvature of both of the transducers was 10 cm and both of them were cut into approximately 1 cm² square elements, as shown in Figure 3. The total number of elements in both arrays was 64 although only 60 were driven in the experiments due to hardware limitations. The ceramic bowl was cut using a diamond wire saw so that the elements were completely separated by a 0.3-0.5 mm kerf. The kerf was filled with silicone rubber that kept the array elements together and isolated them acoustically. The silicone rubber allowed the transducer elements to vibrate with minimum amount of clamping. Each transducer element was connected to a coaxial cable and a matching circuit that was individually tuned. The arrays were similar to the one described in Fan et al, supra, at Figure 1 and the accompanying text, the teachings of which are incorporated herein by reference. The array was driven by an in-house manufactured 64 channel driving system that included an RF amplifier and phase shifter for each channel. The phase and amplitude of the driving signal of each channel was under computer control, as described in Buchanan et al, supra, e.g., at Figure 2 and the accompanying text, the teachings of which are incorporated herein by reference.

In addition to phased arrays configured as described herein and shown in the accompanying drawings, phased arrays can also be constructed in accord with the arrangements described and shown in co-pending, commonly assigned patent application 08/747,033, filed November 8, 1996, the teachings of which are incorporated herein by reference.

Using the methodologies and apparatus described above, it is possible to produce a sharp focus through the skull with the single element transducer when the operation frequency was 1 MHz or less. The beams had secondary peaks introduced by the skull but the main peak was the largest. The location of the peak was shifted by the skull by 1-3 mm from the geometric focus, as shown in Figures 4A - 4H. However, the focus was completely obliterated with an operating frequency of 1.67 MHz, as shown in Figure 4H.

The maximum pressure amplitudes achieved at the maximum power output of the amplifier were frequency dependent and are given in the table below. The maximum average pressure amplitude at the frequency of 0.554 MHz was 8.0+/- 0.6 MPa. Average Pressure P+ P-

Freαuencv fMHz Amplitude (MPa) fMPa (MPa

0.248 3.8 4.2 3.3

0.559 8.0 8.9 7.1

1.0 5.9 6.9 4.5

Figure 5 shows the image across the brain for the first of the sonications and demonstrate tissue damage indicated T2 changes. The tissue damage was also visible in Tl images with and without contrast enhancement.

To measure the phase distortion caused by the skull, a hydrophone was placed in the geometric focus of the array under test. The skull was placed between the array and hydrophone and each transducer element was powered separately in sequence while recording the time difference between the reference signal and the acoustic wave at the focus. This was done with both of the arrays. The phase changes required to correct all of the waves to arrive at the same phase at the focus are plotted in Figure 6.

To investigate the effect of the phase correction the pressure amplitude distributions were measured in water by scanning the needle hydrophone. The main impact of the phase at 0.6 MHz was in the location of the focus which could be corrected back to the geometric focus. This is shown in Figure 7 A, which illustrates the pressure amplitude profile across the focus of the 0.6 MHz phased array in water. Figure 7B shows the pressure amplitude profile across the focus through bone. And, Figure 7C shows the pressure amplitude profile through bone when a phase correction according to the invention is used. Figure 8 likewise illustrates the pressure amplitude distribution along the central axis of the array with and without phased correction. The magnitude was reduced to 33 % and 40 % of its water value without and with the phase correction, respectively. The importance of the phase correction was demonstrated more clearly with the higher frequency array. With this array the focus was completely destroyed by the skull (Figure 9b). However, when phase correction was introduced, the focal spot was returned into its original shape (Figure 9c) with the half-width of the focus of about 1 mm. The insertion of the skull reduced the peak pressure at the focus of the 1.58 MHz array to about 5 % of its water value when phase correction was applied.

The embodiments of the invention discussed above and shown in the drawings provide improved methods and apparatus for neural diagnosis and therapy through application of short, high intensity ultrasound beams that induce cavitation at selected locations within the brain. These and other embodiments can be beneficially used to deliver focused ultrasound beams to the CNS tissues and fluids, thereby, permitting their ablation or other physiological modification. Thus, for example, the embodiments can be used to ablate tumors, cancers and other undesirable tissues in the brain. They can also be used, for example, in connection with the technologies disclosed in copending, commonly-assigned U.S. Patent Application No. 08/711,289 (the teachings of which are incorporated herein by reference) for modification of the blood-brain barrier, e.g., to introduce therapeutic compounds into the brain. Because they do not require that portions of the skull be removed, the embodiments permit the foregoing to be performed noninvasively.

The results also show that adequate ultrasound transmission can be induced through human skull to induce cavitation in vivo. This can be done with single element applicators, e.g., preferably at frequencies less than 1 MHz and at higher frequencies with phased arrays that correct the phase distortion caused by the variable thickness of the skull.

The maximum pressure amplitude of 8 MPa induced through the skull at 0.559 MHz was able to induce cavitation damage in vivo rabbit brain. This value was reached through an area of 10 cm in diameter to a focal spot diameter of about 5 mm (50 % beam diameter). If the whole available skull surface around the brain is utilized, then a window of at least three times larger could be used. In addition, the geometric gain would allow the peak power through the skull to be increased. Acoustic power up to 30- 80 W/cm² of the transducer surface area for continuous wave sonication can be generated by ceramic transducers. Higher peak powers could be achieved with the pulsed sonication used for induction of cavitation. Thus, it is estimated that much higher pressure amplitudes than measured here can be induced in the brain through the skull.

The values measured in connection with the foregoing compare favorably with the 4 MPa that was reported to be the threshold value in vivo muscle at 0.6 MHz and a value of 8.5 MPa at 0.936 MHz in vivo rabbit brain (the threshold at 0.6 MHz would be lower since it has been shown to decrease with frequency). Thus, the results demonstrate that adequate ultrasound transmission through skull can be generated to induce cavitation in the brain.

Our results demonstrate that low frequency beams can be focused through the skull, though, the focus may be shifted from its geometric location. Therefore, it can be helpful to detect the focal spot location in the brain prior to the therapy exposure, e.g., using magnetic resonance imaging to detect the local temperature elevation or cavitation in the brain at exposure conditions that are below the tissue damage threshold. For example, low power test exposures can be delivered through the skull while using MRI to detect the location of the focal spot. Based on the imaging information the location can be corrected to overlap the target volume prior to the therapeutic exposure.

The phase measurements with the arrays support the observation made with the single element transducers showing that at 0.6 MHz 80 % of the phase errors caused by skull are less than 90° and thus, each wave is adding to the pressure wave at the focus. However, at 1.58 MHz over half of the waves had phase shifts that caused the waves to arrive out of phase at the focus. This observation can be explained by the difference in wavelength that is 2.50 mm at 0.6 MHz and 0.95 mm at 1.58 MHz. The possibility of inducing selective thermal damage at the focus, without damaging the skin or brain surface, may be possible due to the small focal spots achieved with the phase correction. However, the thermal exposures have to be short to reduce blood flow and perfusion effects that are strong in brain tissue. The sharp temperature gradients at the focus transport more energy away from the focus than in the bone where the beam is wide and the gradients shallow. At 1.58 MHz, full utilization of the skull surface may provide marginally adequate geometric gains to overcome the skull heating problem. However, at lower frequencies especially around 0.5 MHz the focal brain tissue thermal therapy seems feasible although not as likely as utilization of cavitation effects.

The results demonstrate that the effects of the skull to the beam shape can be eliminated using a phased array with proper phase corrections. In the example above, the phase correction was calculated from hydrophone measurements. The same corrections can be made by measuring the skull thickness from CT or MRI scans and then calculating the phase correction required for each array element. The same may be accomplished by sending a short ultrasound pulse from each or selected elements of the of the phased array and then listening for the echo back from the inner surfaces of the skull or other structures in the brain. The effect of the skull on the wave propagation at each location could then be calculated. This can also be done before therapy by mapping the skull effect using ultrasound.

Although good results were achieved with only 60 transducer elements in the phased array, still more and smaller elements may facilitate moving the focal spot inside of the brain. Similarly, at higher frequencies, smaller elements may allow better phase correction further reducing the losses induced by the skull.

Thus, the geometric gain of about 20 that is required to compensate for the losses caused by the skull can be easily achieved by focusing. This is larger than the gain of 10 required to compensate the average losses. This indicates that adequate power for induction of cavitation can be delivered using phased arrays through the skull even at frequencies that are too high with a single element applicator.

In summary, the invention provides methods and apparatus for noninvasive diagnosis and treatment of the brain using cavitational mechanism and pulsed ultrasound. It permits adequate power transmission through the human skull can be induced to cause tissue damage while keeping the exposures in the overlying tissues below the cavitation threshold.

Those skilled in the art will appreciate that the invention can be applied for purposes of tissue ablation, as well as in other procedures where focussed ultrasound is desired. These include opening the board-brain barrier, activation of therapeutic agents, occlusion of blood vessels, disruption of arteriosclerotic plaques and thrombi, etc. It will also be appreciated that the invention can be applied for treatment of humans, rabbits and other animals. The embodiments discussed above and shown in the drawings are illustrative only. Other embodiments, incorporating substitutions, modifications and other changes therein, fall within the scope of the invention. These include embodiments with transducer arrays of different sizes, shapes and numbers of elements, as well as embodiments with different amplification and driving systems. In view of the foregoing, what I claim is:

Attachment I

to Patent Application for

Methods and Apparatus for Delivery of Noninvasive Ultrasound Brain Therapy Through Intact Skull

Buchanan et al, "The Design and Evaluation of an Intracavitary Ultrasound Phased Array for Hyperthermia," IEEE Trans. Biomed. Eng.. v. 41, pp. 1178 - 1187 (1994)

Design and Experimental Evaluation of an Intracavitary Ultrasound Phased

Array System for Hyperthermia

Mark T. Buchanan and ullervo Hynynen

Abstract — For evaluating the feasibility of treating prostate itations: the depth to which they can effectively heat, and cancer, a 64-element linear ultrasound phased array applicator their limited ability to control the power deposition fields for intracavitary hyperthermia was designed and constructed. [2]. [3 ). Both deeper penetration and increased control over A 64-chaππel ultrasound driving system including amplifiers, phase shifters, an RF power meters was also developed to drive the power deposition pattern can be achieved using linear the array. The design of the array and driving equipment are phased arrays. Since phased arrays can focus their radiated presented, as are the results of acoustical held measurements energy, they theoretically could heal tissues to therapeutic and in vitro perfused phantom studies performed with the array. temperatures deeper ihan nonfocused arrays. The ultrasound Several techniques for heating realistically sized tumor volumes power deposition pattern can be electronically tailored as were also investigated, including single focus scanning and two techniques for producing multiple stationary foci. The results necessary to produce therapeutic temperatures within a desired show that the operation of the array correlated closel with volume. In areas whe e heating would be undesirable, the the theoretical model. When producing a single stationary focus, phased array could lake advantage of destructive interference the array was able to increase tissue temperature by 12°C in to minimize power deposition at those locations. The ease vitro in perfused phantom. With some minor improvements in by which (he power deposition pattern can be electronically array design, intracavitary phased arrays could be evaluated in a clinical environment. altered in real time provides a means for the compensating for varying physiological parameters, paitenl posilioni g. and for minimizing patient discomfort.

I. INTRODUCTION Several lypes of ultrasound phased arrays have been

INTRACAVITARY ultrasound arrays offer an attractive proposed or built for hypεπhenmic purposes. These include means of inducing local hyperthermia in deep-seated tuUmemura and Cain^'s sector-vortex and concentric ring applimors located near body cavities. By locating the radiators as cators |4], |5], the cylindrical section applicator developed by close as possible to the treatment site, problems frequently Ebbini et at. (6). |7J, as well as the tapered array developed encountered with external techniques, such as blockage of by Benke cr eι al. [8|. Each of these arrays is composed of the acoustic window by bone or gas, or simply the inability anywhere from 16 to 64 individual elements and operates at to attain adequate energy penetration, can be avoided. Early frequencies between 500 kHz and 750 kHz. While these arrays results using multielement, nonfocuscd arrays of half-cylinder show significant potential, they ate meant to be used in external transducers operating at 1.6 MHz suggest that such arrays can applications and therefore are unsuitable for intracavitary use be clinically useful in the treatment of prostate cancer ( 1 ). in their reported configurations.

The proximity of the prostate to the rectum wall makes it Previously, a theoretical study on the feasibility of intracava good candidate for heating using intracavitary ultrasound itary phased arrays using half-cylinder elements had been done radiators. Since the prostate is located very near the anus and by Diederich and Hynynen |3). Based on many acoustical and only millimeters away from the rectum wall, an applicator can thermal simulations, it was concluded that a practical array be easily located close to the prostate. The prostate is one of would be composed of 30 half-cylinder elements with 2.5 the most easily accessible tumor sites, and one that affects a mm center-to-center element spacing operating at 500 kHz. large enough population of patients to be potentially clinically The study predicted that the grating lobes formed with this useful. As such, most of the experiments were conducted with array could be negated by using surface cooling. One of the goal of heating the prostate in mind. the major considerations in specifying this design was not

While the nonfocused arrays have shown considerable pothe performance of the array, but the apparent high cost of tential in heating the prostate, they have two primary lim- the amplifiers necessary to drive the system. The desire to

Manuscripi received November |7. 1992: revised Augus_t 9. 199₄. _This minimize the number of amplifiers led to the specification of work was supported by grant number Oi CA 48939 ^"from (be N.iioπal a 30-e)emenι array. Cancer Institute. Initial array designs were based on some recommendations

The authors are with the Department of Radiology. Division or' _MRI. Bπgharn and Women's Hospital, Harvard Medical School. Boston, MA 02115 from the theoretical study and expanded upon in an attempt USA. to construct a practical intracavitary phased array. It became

IEEE tog Number 9406161. apparent that more than 30 amplifiers would be necessary for

00i«-9294Λ)-l$04 00 © 1994 IEEE BUCHANAN AND HY^NYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

Fig I The 64-etement _rr,ιy ond a single element showing Ihe geometry ot the ,ιtτ_y The urray was mounted to a brass shell 19 mm in diameter The bdik of the applicator *aι removable so that wires could be connected to the inner electrode

a useful array This led to the development of the 64-channcl of commercially available equipment to drive the array, a dπvmg system to overcome the limitations imposed by the 64-channel amplifier system was developed. This system is number of amplifiers This allowed for the evolution of a 64- composed of phase shtfters/duty cycle controllers, amplifiers, channel phased array with 1 73 mm cemer-to-center spacing RF power meters, and impedance matching networks (see and a total array length of 1 105 mm operating at 500 kHz block diagram in Fie 2) and is controlled by an IBM PC compatible computer via a digital I/O interface

U. MATERIALS AND METHODS The amplifiers are based on a switching MOSFET design and are designed around International Rectifier's IR2110 dual

A Array Consπ itction MOSFET driver Each of the amplifiers is capable of deliv¬

A 64-element array was constructed using half-cylinder ering up to 16 of RF power at 500 kHz into a 50-W load transducers operating in their resonant radial mode at 500 kHz from each channel or a total output power of about 850 W. The array was made by slicing washer-shaped elements with The amplifiers convert digital logic input signals into high a diamond wire saw (Laser Technologies. North Hollywood, power sine waves while preserving the phase of the input CA) from 15 mm O D by 30-mro long cylinders of P2T- Nignal The amplitude of the output signal is controlled by 4 material (EDO. Salt Lake City, UT) with silver electrodes two methods All of the amplifier outputs can be controlled plated on both the inner and outer wall sur aces The transducer simultaneously by adjusting the output voltage on a 1000 W slices were glued together using a silicone adhesive (Dow DC supply, or the output of each channel can be individually Corning, Midland, Ml) with 0 17-mm thick silicon rubber controlled by varying the duty cycle of the input signal. The spacers (SPC Technology, Chicago. IL) between each element result of reducing the duty cycle of the input signal (the duty (producing a dead space of 0.23 mm between elements). The cycle is the percent of "on" time of the input signal per clock stack of elements was then cut in half along the axis of the cycle) ts a corresponding decrease in the amplitude of the cylinder and the two half-cylinder sections glued together to output signal. form the full array The array was bonded to a brass shell Since the amplifiers require digital input signals, the phase to form the complete applicator, as shown in Fig. 1 Wires shifting and duty-cycle control is implemented using digital were soldered to the inside wall of each array clement that ,ouπters ( 10], [ 1 1]. These circuits provide 22.5° phase shift extended the length of the shell to the handle where they were resolution from 0-360°. and 16 steps of duty cycle (amplitude) each connected to a 2-m RG-178 coaxial cable The electrodes that allows the output to vary from 0-100% of the output on the outer surface of the array were electrically connected by power, as allowed by the voltage from the DC supply. While single wires embedded in beads of silver epoxy (Chomeπcs, the phase shift resolution may seem somewhat limited, studies Wobum, MA) running along both edges of the array. have shown that it is sufficient for the uses described here [11], [12]

B. Driving Hardware The actual RF output power from each channel is monitored

One of the primary disadvantages of phased arrays is the by power meters that measure both the forward and reflected increased complexity of the driving equipment. Due to a lack RF power [13] The power meters, which were also designed IEEE TRANSACTIONS ON BIO EDfCAU'ENifc. .eERINOrVOL -tf. NO ^,"12.TJE t BER 1994

(b) ig 2. (a) Block diagram of the

dru ms system (h) A Mmplificd MIKΠ I ol the switching amplifiers

TABLE I

and constructed in house, allow for v erification of the output characterized using a hydrophone and the results used to power on each of the amplifier channels, and allow for compensate the excitation signals. the monitoring of the outputs vo that possibly dangerous fault conditions do not occur. B> monitoring the amount of reflected power, changes in the electrical matching and the C Acoustic Modeling aπay operation can b easily monitored so that peak array The 3-D ultrasound modeling program used as originally performance can be maintained developed by Diedench and Hynynen [3] and is only briefly

Finally, to maximize pov*er traile between the amplifiers introduced here. The routine models JV cylindrical radiators of and aπay elements, a simple LC :nι>.d&nce matching network finite length, radius and separation as each of the ultrasound was used on each channel between tne power meters and the transducers in the array, and assumes a target media having array. This not only maximizes p-y->tι transfer, but ensures similar acoustical properties as tissue (Table I). The routine greater unifomuty in the acoustics: output signals for a given models the surface o each of the cylindrical elements as an excitation phase and amplitude ..r.ήcα experimentally by a evenly spaced grid of simple hemispherical sources and uses needle hydrophone; Without eiec-r.cai matching, the phase Huygens principle to sum the contributions of each source at shift and amplitude distortion c-._>ed b> the impedance difeach point in the field All of the simulations were done with ferences betw een elements in τ_*.t -~T- would have to be the sources spaced on a 1/32 grid. BUCHANAN AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

The acoustical pressure field was calculated in the z-plane The technique solves the Rayleigh-Sommerfeld inte ral f using

where where /> is the density of the propagation medium, c is the p i _r = r. ".t " _cl( πf - 2π j. - φ) - ar speed ol sound in the medium, k is the wavenumber, S' is the

(2) surface of the source, _n is the particle velocity normal to thr surface of the source, and r_m- r_n is the distance between a and n_z is the number of sources in the 2-dιrectιon, n_$ is the control point (r_m) and the surface of a source (r_n), m = i, number of sources in the theta direction, P_s is the pressure

2 M, and π = 1, 2» .,., JV. from a single source (Pa), P_sa is the pressure amplitude at the

To simplify programming of this algorithm, (4) can be surface of the source (Pa), r„ is the radius of the source (m), r described by is the radial distance from the center of the source (m), I is the wavelength (m) ώ is the phase of the excitation signal (rad), ϊu = p (5) / is the operating frequency (Hz), and α is the attenuation coefficient (Np/m) where u is the excitation vector, p is the complex pressure at the control points, and H is the remainder of (4) For intracavitary hyperthermia uses, the number of elements in

D Foc using Techniques the array (JV) is always greater than the number of control

The single focus case is the simplest form of focusing that points (M) This leads to an underdetermined system of can be done with a phased array The single focus is produced equations with an infinite number of solutions The minimum by setting the phases of the driving signals so that constructive norm solution (ϋ) can be determined by using a least squares interference occurs at the desired focal position The phase of approximation of (5) each of the driving signals for an array with N elements can be calculated from the differences in the path lengths <f, between u - H'^HH'*) 'j (6) each array element and the focal position by where II" is the complex conjugate of H With this technique, all of the array elements contribute to all of the foci, unlike

= 27r— 1- 2πτ;ι (3) the split focusing technique where each element contributes to only one focus where /, is the phase (radians) of the ι^th element, ( is ihe wavelength (m) t = 1 , 2, , V and rn = 0, 1 , 2, The size III. EXPERIMENTAL TECHNIQUES of the focus produced with this technique is usually too small to heat an entire tumor volume, and therefore other techniques A Acoustical Efficiency Measurements must be used to heat larger volumes

The acoustical output of the cylindrical transducer elements

To heat larger volumes, two techniques were investigated was measured using a radiation force technique [16] The single-focus scanning and multiple focusing The single focus array was placed in a brass cone with 45° sides to reflect scanning routine simply stepped a single focus back and forth the radially emitted ultrasound fields down into the acoustical along a predetermined length of the array The focal depth was absorber [ 17] The force on the absorber was measured using kept constant while the focus was stepped every 300 ms to the a Metier AE 160 (Hightstown, NJ) microbalance, while the next position in the scan Delivered power was maximal at the RF power was measured using a Hewlett Packard 438A RF edges of the scan but was reduced to 64% of maximum power power meter and a Werlatone C2625 (Brewster, NY) dual at the center of the scan to flatten the temperature distribution directional coupler The efficiency was calculated as the ratio in the perfused phantom experiments (the power distribution of the acoustical power to the RF electrical input power. was experimentally determined)

The other technique, multiple focusing, simultaneously produces more than one focus within the target volume The drivB Ultrasound Field Measurements ing signals necessary to produce multiple foci were calculated The ultrasound fields were mapped in a tank of degassed, using two techniques split focusing and the pseudo-inverse deionized water by mechanically scanning a thermocouple To create multiple foci with the split focusing technique, the embedded in a small (2-mm diameter) plastic sphere The ther^¬ array was divided into subarrays, each of which produce a mocouple was positioned by a three-axis computer controlled single focus m the same manner as previously descnbed The scanning table The applicator was mounted on a rotational pseudo-inverse method, developed by Ebbini and Cam [15], device that allowed measurements to be made in a radial arc uses a series of control points that represent the magnitude around the array by rotating the array Measurements were of the ultrasound field at given points A brief summary of made on a J x 1 mm gnd with the recorded data being the Ebbini and Cain's technique follows. average of three consecutive measurements C. In Vitro Kidney Experiments

Alcohol fixed canine kidneys were used as phantoms for studying the heating characteristics of the array. The kidneys had previously been prepared as described by Holmes et al. [18], and were rehydrated prior to use. The experiments were conducted at room temperature using degassed, deionized water as the perfusate. A metering pump (Fluid-Metering Inc., RHICKC, Oyster Bay, NY) connected to the renal artery circulated water through the kidney while the renal vein was allowed to drain into the tank. The kidney was held in place by gently sandwiching it between two PVC membranes mounted Pu σ

to a Plexiglas frame. The applicator was firmly clamped to the IΛfatar Bach frame to maintain a fixed distance between the surfaces of the Fig. 3. Diagram of ilie in vitro kidney experimental setup. kidney and array. Fig. 3 shows a diagram of the experimental setup.

Temperatures were measured in the kidney using either 20 - seven sensor probes sutured in place perpendicular to the array through the focal region, or by one or two single sensor thermocouples pulled along a path parallel to the array. In experiments conducted with the multiple sensor probes, the kidney was exposed to ultrasound for a total of 10 min with temperature measurements occurring every 30 s. During each temperature measurement, power to the array was disrupted for approximately four seconds (one second prior to the first reading, and up to three seconds to read all of the thermocouples). Since this technique does not give very fine spatial detail of the temperature distributions, the pull-back

technique was more frequently used.

The pull-back experiments were conducted by pulling one or two single uncoated thermocouples (0.05-mm w ire) by a 10 15 20 computer controlled stepper motor along a track parallel to Electrical Power (W) the array through the kidney tissue. The temperatures were Fig. 4. Acoustical output power as a function of electrical input power for measured every 1 mm and the average of three readings was a single 30-mm. full cylinder element, and arrays made up of 16 2.2-mm and recorded. Prior to each of the experiments, a baseline tem1.5- nι elements with a 0.3-mm insulator between elcmenis. perature was established along the path of the thermocouples. The kidney was exposed to ultrasound for 20 min to allow (width < wavelength), which would produce nearly spherical the temperatures to reach steady-state before the temperature wavefronts. The collimation is necessary to assure the waves profiles were measured, The difference between the baseline reflected by the reflector are normally incident to the surface and final measurement was used to calculate the temperature of the absorber. rise. The kidney was allowed to cool to room temperature The results of these experiments are shown in Fig. 4, None (typically, 30 min) before to the next experiment was started. of the arrays exhibited very high efficiencies despite the low

The thermocouples were mainly located in the medulla of operating frequency. Though not shown here, some results the kidney. The steady-state temperatures in the medulla have were verified using calorimetric techniques. While the 30-mm been shown to be a strong function of the flow into the kidney long full cylinder element had an efficiency of 71 %, the arrays [19]. The flow values were kept relatively low in order to exhibited efficiencies of 27% and 17% for the 2.2-mm and 1.5- simulate the perfusion in the prostate [20], [21 ]. m arrays, respectively. These efficiencies indicate that about 80% of the electrical power delivered to the 1.5-mm wide

IV. RESULTS elements used in the phased arrays is lost either as heat in the transducer or in the electrical matching and transmission lines between the amplifier and the array.

A. Acoustical Efficiency Measurements

The acoustical efficiency of the 500-kHz PZT-4 transducer

B. Ultrasound Field Measurements material was measured for a full cylinder, 30 mm long, and for two half-cylinder arrays with 2.5 mm and l .8-mm center- The effects of element spacing were studied by measuring to-center element spacing (2.2 mm and 1.5 mm element the ultrasound field distributions produced by two 16-element widths) with total array lengths of 24,7 mm and 28.5 mm. arrays with 1.8-mm and 2.5-mm center-to-center element respectively. Small arrays were tested so that the resulting spacing. Fig. 5(a) and (b) shows the acoustical field plots made field would be collimated, unlike single array element in a plane parallel to the array and normal to the surface of BUCHAS -Λ AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

(a) (b)

Radial Distance (mm) Radial Distance (mm)

(0

Fig. 5. Field plots in water of two lo-elemem arrays with l ft-mm (a) and 2 5-mm (b) element center-to-cemor spacing and their corresponding simulation results (to and (d)> Both arrays were focused 30 mm troπi the .iπay surt-cc and 15 mm from the central axis

the half cylinder elements for t o 1 -element arrays with 1 8- So far. all of the acoustical field measurements presented mm and 2.5-mm element spacing along with their respective have been in a plane parallel to the length of the array and simulated fields (Fig. 5(c) and (d)). Both arrays were focused normal to the surface of the array. Ideally, the acoustical field 30 mm from the array surface and 15 mm above the central would taper off toward the edges of the array but be more or axis of the array. The most significant effect illustrated here is less uniform around the arc of the array Unfortunately, this the generation of the grating lobe bv the 16-element array with ideal docs not match reality, as can be seen in Fig. 7. The 2.5-mm center-to-center element spacing (Fig. 5(b) and (d)) acoustic field was measured as a function of rotation angle Since such a large grating lobe not only reduces the power along a line by fixing the position of the thermocouple in of the focus but also could cause heating in unwanted areas, the focal region and rotating the array about its axis. The arrays must be designed to minimize grating lobe formation. acoustical intensity is not at all uniform and, in fact, vanes The array with 1.8-mm center-to-center spacing produced as much as 50% before tapering off at the edges. All of the much belter field distributions than the array with 2.5-mm 2-D acoustical field plots shown here were made at the 0° center-to-center spacing. Therefore, in order to have center- rotation angle where the intensity is only about 50% of the to-center spacing 1 8 mm or smaller, 1.5-mm wide elements peak. Measurements made, but not shown here, show that this were used in the final array design and the dead space between fluctuation in pressure amplitude due to the rotation angle only elements was reduced to 0.23 mm alters the peak intensity and does not effect the overall shape

The complete measured and simulated plots of the ultraof the acoustic field. sound field produced by the 64-element array with 1 73-mm Fig 8(a) and (b) shows the acoustical field plots produced element spacing are shown in Fig. 6(a) and (b), respectively. using the two multiple focus techniques: pseudo-inverse and The array was focused 40 mm deep, and 30 mm from the split focusing. Below them, in Fig. 8(c) and (d), are their central axis. Note that only a small grating lobe is produced respective simulation results. Both techniques were used to by this array even though it is focused a substantial distance synthesize foci 40 mm from the array surface and 10 mm from its central axis. In fact, this array is capable of focusing to on either side of the central axis of the array (20 mm apart) about 35° from the central axis of the array without producing using the 64-element array. Notice that the pseudo-inverse significant grating lobes. technique produces sharper foci than those produced by the IEEE TRANSACTIONS ON BIOMEPIC^I, frfG'lr- "YlP 4l„INβl 11. DBCEMBBΛwl

Radial Distance (mm) Radial Distance (mm)

(*) <°)

Fig 6 Field plot of the 64-element array (a) along with the simulation results (b) for focus 40 mm Irom the turfjce of ihe array and 25 mm off its central axis

10 —j Λ the circulating water in the water bath All of the following

03 —' remit were obtained in the center of the kidney where the focus was located and the surface effects were smallest. The OB -j temperature rises achieved in the middle of the kidney indicate 07 that therapeutic temperatures can be achieved with a stationary -fl 1 single focus at realistic perfusion levels Typical cross-section

profiles measured in the middle of the kidney with the pull- 05 - back technique for a center focus and 15 mm on either side

04 —, are shown in Fig 9(b)

The three different techniques for heating larger volumes: 03 — pseudo-inverse, split focusing, and scanned single foci arc 0 2 compared in Fig 10(a) and (b) The two multiple focusing

01 techniques were used to produce foci 40 mm from the surface of the array and 10 mm on either side of the array's central 00 ^"I^" axis, white the scanning routine produced a series of single

90 -7S .60 45 -30 15 0 15 SO 45 60 75 90 foci 40 mm from the array surface with a total span of Rotational Angle (Degrees) 20 mm The thermocouples were located 8 mm and 15 mm from the surface of the kidney nearest to the array

Fig 7 Single dimensional radial held plot made within the focus as a function of the ^ngle of rotation of the array The center of the arrav arc (35 mm and 42 mm from the surface of the array). The is defined as the zero point scanning technique produced a narrower profile than the two multiple focusing techniques and the distribution lacks the temperature drop between foci The two multiple focusing split focusing method This occurs because the pseudo-i crse techniques produce virtually identical profiles, although the technique utilizes the entire length of the array allowing it to pseudo-inverse technique does not produce as large of a produce sharper foci than the spin focusing technique which, temperature πse on the deeper thermocouple (beyond the since it divides the array into two subarray , effectively uses locus) as does the spin focusing technique. This IS caused b> array lengths only half as long as the actual length of the array. the increased sharpness of the foci produced by the pseudo- inverse technique, as previously explained.

C In Vitio Kidney Exper iments

Fig. 9(a) shows the temperature πse ver us time along a fixed seven-sensor thermocouple probe located perpendicular V DISCUSSION AND SUMMARY to the array The distances marked denote the distance of An intracavitary ultrasound phased array composed of half- the thermocouple from the edge of the kidney nearest to cylinder transducer elements has been constructed for inducing the array. The array, with 32 active elements, was focused hyperthermia in the prostate via the rectum. A 64-channel along its central axis and 10 mm into the kidney (Thirty- amplifier system has also been designed and constructed to two active elements were used since the array length in this drive the phased array. As was shown by the acoustical field configuration could adequately cover the kidney ) The kidney plots, the array is capable of producing focused fields as well as was perfused at a rate of 2.9 kg m^-J s~^l Note that the kidney fields containing more than one focus The arrav is currently is only 35 mm thick and thus, the temperatures close to the capable of producing a 12°C temperature πse in a perfused surfaces of the kidney are dominated by cooling caused by phantom using a stationary focus, and smaller temperature BUCHANAN AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

Radial Distance (mm) Radial Dbttaπca (mm) (e) (d)

Fig. 8 Acoustical field plots produced u»ιne the pseudo-mierse (a) and split focusinj (b) techniques and the related simulated results ((c) and (d)). The foci were produced 40 mm deep, and 10 mm on either side of the central axis of the arr-y

increases were achieved in larger volumes by scanning the expenment because the impedance of the array elements did focus or by creating multiple stationary foci This was achieved not return to their oπginal impedance after cooling. The with flows to the kidney that should simulate the relatively low impedance shifts made it difficult to accurately control the perfusion in prostate [20], [21] radiated power delivered to the kidney Other problems with

What has yet to be conclusively shown is such an array's the array included arcing between elements and water seepage ability to heat tissues to therapeutic temperatures in a regubehind the array that reduced the array efficiency in subsequent lar and predictable manner. The peak temperatures achieved experiments until the array was repaired and rematched. With within the perfused phantoms tended to vary considerably from a more carefully constructed array, most of the power limiting experiment to experiment, making direct comparisons between problems experienced can be avoided the absolute temperature profiles difficult Therefore, the temThe acoustical field plots were m good agreement with the perature profiles that compare various focusing techniques theory, and the techniques for heating larger volumes were were normalized to better illustrate the overall differences all functional, though the multiple focusing techniques were in the shape of the temperature profiles. The fluctuations in the most effective Overall, phased arrays show considerable temperature were caused by a variety of problems, including potential for improvement over currently used intracavitary morphological differences in kidneys and the location of the ultrasound hyperthermia system. thermocouples within the kidneys, the inefficiency of the array, While phased arrays allow significantly more control over and the open loop manner in which the array was ooerated the acoustical field, the current design using half-cylinder

The electrical efficiency of the array would usually drop radiators still lacks control in the angular direction (around the considerably during the first experiment, and dunng each subarc of the array). Additionally, since the cylindrical radiators sequent experiment due to changes in the electrical impedance do not have uniform angular intensities, the angular heating of the array elements. Two primary factors were responsible pattern is somewhat degraded, though thermal conducuon will for the observed changes in the electrical impedance of the probably smooth the resulting temperature distribution. Similar array elements: The first was a thermally induced impedance fluctuations in the angular field distributions have been shown drift caused by the array self-heating during sonication. The for other cylindrical transducers [2], [11], Radial control second impedance shift that occurred from experiment to could be achieved by dividing the half-cylinder elements into IEEE TRANSACTIONS ON BIOMEDICAL ENCIN*^ftRlNG. VOL. 41. NO 12 DECEMBER 1994

Ailil Distance (mm) Fig 10 Normalized temperature/power profiles measured in a perfused

(b) kidney for multiple focusing using the pseudo-inverse and spin focusing techniques as well as single-focus scanning The multiple foci were created

Fig 9. Temperature profile measured in perfused kidney u ing 32 active 40 mm deep. 10 mm on cilher side of the ceniral axis while ihe scanning elements focused 30 mm from the array surface while the kidney a perfused was done with a single focus 40 mm deep and scanned 10 mm on either al the rate of 2.9 kg m^_1 s" ^l (5 mL. mm" ^l flow) The array as positioned side of the central HIS The measurements were made 8 mm (a) and 15 mm 18 mm from the kidney surface The thermocouple locations indicate ihe (b) from the surface of ihe kidney with the array positioned 27 mm from the distance from the surface of Ihe kidney (a) The axial distribution measured kidney surface 12 mm from the surface of the kidney using a pull-back thermocouple (b) in a separate expcπment minimized grating lobe formation, higher frequencies would pie-shaped subelements and driving each subelement indepenincrease the power absorption in tissues Since most tumor dently. However, this would not only dramatically increase sites for which this applicator would be used occur near the number of amplifiers, but the resulting subelements would the cavity wall, the deep penetration allowed by the 500- have virtually no inner electrode due to the wall thickness kHz operating frequency is not necessary. The difficulty in of the element This would senously degrade the already low increasing the operating frequency is that the element size and acoustical efficiency of the cylindrical radiators, center-to-center element spacing would need to decrease as

A possibly better array design would utilize planer array the frequencv increased, making atfay design more difficult elements mounted on a rotating platform By tilting the array As a conclusion, the intracavitary, electrically focused array back and forth, the same volume could be treated as with did demonstrate, at practical depths for prostate treatments the cylindrical radiators. By controlling the acoustic field and (about 30 mm), the feasibility of using phased arrays for output power as a function of tilt angle, 2-D control over intracavitary hyperthermia purposes With a more carefully the power deposition field can be achieved without requiring constructed array, most of the power limiting problems ex^¬ additional amplifiers. An additional improvement on the array perienced can be avoided The acoustical field plots were design would be to increase the operating frequency. While the in good agreement with the theory, and the techniques for 500-kHz operating frequency allowed larger element sizes and heating larger volumes were all functional, though the multⁱple BUCHANAN AND HYNYNEN INTRACAv rARY ULTRASOUND PHASED ARRAY SYSTEM

Attachment II

to Patent Application for

Methods and Apparatus for Delivery of Noninvasive Ultrasound Brain Therapy Through Intact Skull

Fan et al, "Control of the Necrosed Tissue Volume During Noninvasive Ultrasound Surgery Using a 16-Element Phased Array," Med. Phvs.. v. 22, pp. 297-308 (1995)

Control of the necrosed tissue volume during noninvasive ultrasound surgery using a 16-element phased array

Xiaobing Fan^a) and Kullervo Hynynen

Department of Radiolog^y, Brigham and Women * Hospital Harvard Medical Sthool Boston. Massachusetts 02115

(Received 24 March 1994; accepted for publication 31 October 1994)

Focused high-power ultrasound beams are well suited for noninvasive local destruction of deep target volumes In order to avoid cavitation and to utilize only thermal tissue damage, ,hιgh frequencies (1-5 MHz) are used ultrasonic surgery However the focal spots generated by sharply focused transducers become so small that only small tumors can be treated in a reasonable time. Phased array ultrasound transducers can be employed to electronically scan a focal spot or ιo produce multiple foci in the desired region to increase the treated volume. In this article, theoretical and expenmental studies of spherically curved square-element phased arrays for use in ultrasonic surgery were performed The simulation results were compared with expenmental results from a 16-element array It was shown that the phased array could control the necrosed tissue volume by using closely spaced multiple foci The phased array can also be used to enlarge a necrosed tissue volume in only one direction at a time, i e . lateral or longitudinal. The spherically curved 16 square-element phased array can produce useful results by varying the phase and amplitude setting Four focal points can be easily generated with a distance of two or four wavelengths between the two closest peaks The maximum necrosed tissue volume generated by the array can be up to sixteen times the volume induced by a similar spherical transducer Therefore the treatment time could be reduced compared wuh single transducer treatment.

Key words phased array, ultrasonic surgery I. INTRODUCTION dred. The main disadvantage associated with using a large number of elements is that the same number of amplifiers

In ultrasound surgery, to avoid cavuaπon and have a sharp and electronic circuits are also required boundary between the tumor and normal tissue, high freThe motivation of this article was to demonstrate that a quency focused transducers have to be used ' The focal spots phased array system wπh a small number of elements can be generated by sharply focused ultrasound transducers are utilized to provide control over the focal spot size for surgismall when compared with the diameter of many tumors cal purposes. The excitation signals of the elements vary in Obviously this type of focus is not efficient to treat large both amplitude and phase Hence the major task in utilizing tumors, which would require a large number of exposures to phased arrays is to determine the phases and amplitudes cover the whole target volume. However, in order to accu needed to produce desired ultrasound fields. The inverse ra^ly treat the target volume close to cπtical structures, technique introduced by Ebbini and Cain⁶ can be used for s focal spots may be required dunng pan of the therapy direct synthesis of ultrasound fields with multiple foci. SevThus, controllable focal spot size is required. The most ateral amplitude and phase settings were calculated for differtractive method to obtain control over the focal spot size is to ent sets of selected control points. The limitation of the numuse phased arrays. ber of utilizable control points was also investigated for a

Phased array applicators were introduced to ultrasound given array. Several simplified amplitude and phase settings hyperthermia cancer therapy in the early 1980's. Dunng the based on the calculated amplitudes and phases were empast decade, many efforts have been made to investigate the ployed for ultrasound field calculations. These dπving signal advantages of phased arrays in hyperthermia. and several sets can be utilized, when different focal spot sizes are rephased array applicators have been developed. Phased array quired for the array proposed hexe. The transient bioheat applicators can be divided into the following categones: antransfer equation was employed to estimate the temperature nular or cσncentπc-πng arrays,^2-3 stacked linear-phased elevation due to the ultrasound power deposition. Then the arrays,'' sector-vortex arrays,³ tapered linear-phased arrays,⁵ necrosed tissue volume was predicted by the isothermal dose cylindπcal-section arrays,⁶ and square-element spherical- volume Computer programs were also used to do a paramet- section arrays.⁷ It was shown that ultrasound phased arrays nc study to investigate the influence of the dimensions and can provide good control over the heating pattern with flexfrequency of the array on the necrosed tissue volume. ibility in moving the focus and producing multiple foci Previous research in ultrasound phased arrays has been concenII. MATERIALS AND METHODS trated in hyperthermia cancer therapy Relatively low A. Square-element array with spherical surface frequencies (about 0.5 MHz) and small element size (compared to the wavelength) were employed in these studies The phased array was made from a spherically curved The studies on square-element sphencal-section arrays emtransducer which was divided into small square elements phasized using a large number of elements, up to a few hunThe whole array was airbacked. It was constructed from a

297 Med. Phys. 22 (3), March 1995 0094-2405/9522(3^287/10 S1.20 © 1995 Am. AΘΘOC. Phys. Med. 297 298 X. Fan end K. Hynynen: Control of necrosed tissue volume during ultrasound surgery 298 where j = J-T, ύ is the density of medium, c is the .peed of sound, k is the real wave number. o„ = Λ .β^J<ω'⁺ '^,,) is the complex excitation source of the nth element with amplitude A„ and phase θ„ , r_m- is the distance from :ι point (x„ ,y_n ,z_r) on ihe nth element to the field point of interest (x_m ,y_m ,z_m), S„ is the area of the nth element.

In ultrasonic surgery, water is used as a coupling medium between the transducer and tissue interface. Water is considered a nonattenuatiπg'medium. and tissue a weakly attenuating medium. The effects of attenuation can be satisfactorily described by replacing e ~^ji'"'^π with e~^lk<^rm« in expression

(1), where k_c ~ k -ja is the complex wave number with attenuation coefficient α. " ¹²

The power deposition for the desired volume is given by

FIG. I. Configuration of the 'quar -element spherically curved phased array with its coordinate system,

spherically curved PZT4 bowl, having a diameter of 100 where /-_abs is the absorption coefficient and Z=pc is the mm. a radius of curvature of 130 m , and a frequency of 1.4 impedance of the medium. MHz The bowl was cut into 16 elements, each with a length of 20 mm per side A 0.3-mm space between the elements was filled with silicone rubber for electrical and mechanical D. Inverse technique isolation. Each of the elements was connected to an LC matching circuit to match the impedance to 50 i and 0° The The inverse technique can be used to calculate the ampliarray was driven \* ιth a custom made 16 channel amplifier tude and phase settings from selected control points where (Labihermics, Champaign. Illinois). The phase and amplithe desired pressure values are given. Defining tude were controlled by RF signals feeding the amplifiers ;ρcit f _e ~^Jkr»« The driving signals were generated by an in-house manufach ~ — dS„, (3) tured digital circuit ' The amplitude and phase of each input signal were digitally controlled with 3-bit resolution in amthen for M field points, expression (1 ) can be written in plitude and 4-bit resolution in phase The number of ele matnx notation as ments in each row and each column were the same so that P = HU, (4) the whole applicator was a square shaped focused transducer The configuration of the phased array for the expeπments is where shown in Fis 1

B. Ultrasound field measurements

The relative pressure amplitude squared distributions

^TMNI were measured in degassed water using a needle hydrophone (active spot size 1 mm) scanned across the focal region. The needle hydrophone u as moved by stepper motors, typically with 0 I -mm steps across the beam The total acoustic power was measured using a radiation force technique.

"N

The elements of matrix H are evaluated by numerical inte¬

C. Computation of ultrasonic fields gration using expression (3). Equation (4) has two important features. First, if the H matnx is calculated and saved for a

Consider an ultrasound phased array with N elements, desired field, then the pressure field can be evaluated using employed to produce an ultrasonic field in a πonattenuating Eq. (4) instead of Eq. (1 ) for a given excitation source matrix medium Assume the coordinate system is defined as in Fig. U. Computation time can be saved by using this method 1. According to the theory of ultrasonic radiators developed when varying the excitation source over the same calculation by O'Neii"¹ and the principle of superposition of acoustic volume. Second, for a desired pressure field pattern, i.e., for pressure, the pressure field due to this transducer can be a given matnx P, the excitation source U can be calculated evaluated by the ^' Ra> leιgh-Sommerfeld integral The inte- by eral is υ = H*p where W⁴ is the pseudomvcrse matrix of H. Usually the total

number of control points are much less than the total number TABLE 1. Selected control poinit use for imeme ca ul. Jiv for 16 square- where T is the temperature iiat |ιjne||. 1 HMJP^cat«» ■»*^'< « element spherical curved phased array The conlrol points are located al p, is the density of the tissue, c, is the specific heat of the ;= 129 mm plane Note the unirs for ι and v are millimeter;. tissue, k, is the thermal conductivity of ihe tissue, w is the

.\ Phase blood perfusion rat, . c,, is tl... specific heal of the bloc.l. T„ is the arterial blood temperature, and Q(x.y.;) is the acous¬

1 25 0.0 0.0' tic power deposition rate per unit mass. The thermal response

Case 0,0 ! 25 0.0' 1 ^■ 1 25 ob 0.0' was simulated in a homogeneous medium. A surface tem¬

0 0 - ) 25 o.rr perature of 37 ^CC on the cube and an initial temperature of 37 °C inside the cube were used as boundary and initial con¬

1 ,25 0 0 00°

0 0 ditions for all of the computations. A numcπcai finite differ¬

Case 1.25 90 0' II ^■ 1.25 0.0 1800* ence rneihod was employed to solve Eq. (5). Previous studies

0.0 - 1.25 270.0' have shown that there are several parameters which affect the temperature elevation.'^-1 Temperature elevation has been

3 25 0.0 0.0°

Case 0.0 3 25 90.0° shown to be almost independent of the blood perfusion rate III ^•3 25 00 180 0' for short ultrasound pulses. '* For this reason, ultrasound

0.0 -3 25 2700= pulse durations of 1, 5, ard 10 s were used in this study. The maximum temperature reached in all simulations was kept

5.25 0 0 00^<

Cane 0.0 5.25 90.0' the same (80 °C) by adjusting the input power IV ^• 5.25 0 0 IS0.0'

0 0 -5 25 270 0"

F. Thermal dose calculation and necrosed tissue

3 25 0 0 0 0' volume estimation

1.62 2 51 60.0°

Case ^■ 1 62 2 51 120.0° The thermal dose calculation was based on the technique v -3.25 00 180.0" suggested by Sapareto and Dewey '⁵ Using this technique,

- 1 62 - 2 81 240 0° the accumulated thermal dose was calculated at a reference

1 62 - 2.81 3000^C temperature by numerical integration under different tem¬

3.25 0 0 0.0° perature profiles The thermal dose, i.e., equivalent time, at

2.3 2 3 .15.0° the reference temperature can be evaluated by

0.0 3 25 90.0°

Case -2.3 2 3 135 0° '^β 'l|«.l V1 - 3 25 0 0 1800° Dose(; _τ ) = \"'*°* _Rτ_κr- πn _{dt ==} 2 ^* _Λr,_tf-r_Λ,_Δr>

-2 3 -2.3 225.0' J ι=0

0 0 - 3.25 270.0° (6)

2 3 -2.3 315.0° where T„₍ is the reference temperature. ι_fin3,= r_hβ,„_ng+f_Cooiι-_>_ is the final time, Δ; is a small time interval. T_Λf is the aver- as;e temperature during time Δr. and R is a parameter given of excitation sources in this case. The H^τ matrix can be by ii'-rl b 0.5. if r(r)Ss 3 control points were selected in a plane at a desired R = focal position. All the points were evenly distributed on a 0.25 , otherwise circle. Since the pressure at a control point is a complex The necrosed tissue volume is estimated by the isothermal number, both the magnitude and the phase are required to dose volume surrounded for 240 mm at the reference temperform the inverse calculations. In selecting control points perature of 43 °C. This technique has been found to be a in this manner, it is natural to make the magnitudes the same reasonable model for predicting tissue necroses induced by a for all the points, and the phases either in phase or out of spherically curved transducer.¹" phase so that the phase rotates around the central axis. The A two layered medium water-tissue was assumed in the latter setup gives destructive interference on the central axis simulations. The speed of sound and the densify were 1500 thus, eliminating potential hot spots on the axis.³ The sem/s and 998 kg/πv\ respectively, for both media. The attenulected control points used in the inverse calculations are ation coefficient of the tissue was assumed to be 10 Np/m given in Table I. MHz. The thermal properties of the tissue are given in Table II.

E. Thermal modeling

An approximate temperature response to the power depoIII. RESULTS sition was predicted by the transient bioheat transfer equaA. Comparison of experimental results with tion. The differential equation is simulations

It was necessary to veπfy the numerical model using the ^p'.^c' 16 square-element phased array with a spheπcal surface be

fore relying on the simulations. The experimental and si u- edleal Physics, Vol. 22, No. 3, March 1995 aoυ Λ. ran ana

T .BLE II. Thermal properties of ihe tissue used in >he simulations. TABLE III. Phase and amplitude setungs calculated with the inverse technique for the 16 square-element spherically curved phased array.

Parameter Value Units

Λ, p, — tissue density 998 kg m' c, — tissue sper .. heat 3770 J kg *C — r ₇. 64 2° 64.2° 265.7" 1.0 2 6 2.6 1.0 λ,— tissue thermal conductivity 0.5 W/m °C C.se 64 2° 58 1 ° 5f ' ^• 64.2* 2.6 6.2 6.2 2.6 w — blood perfusion rale kg/m' s I 64 2° 58.1 ° 58.1° 64 2* 2.6 6.2 6 2 2.6 c_h — blood specific heat 3770 J/ g 'C 265 7° 64.2° 64 2° 265.7° 1.0 2 6 2.6 1.0

T_a — πenal blood temperature 37 'C

119 7^* 95.0°, 38 0° 29.7° 1,8 1 5 1.5 1 8

Case 128 0" 104.1* 14.1° 5.0° 1.5 1 0 i.o 1.5 II 185.0° 194 ir 284.1° 3O8.0" 1.5 1.0 1.0 1.5 lated profiles of the pressure amplitude squared at the acous209.7° 218.0° 275.0° 2997° 1.8 1 5 1.5 1.8 tic focus are displayed in Fig. 2. All the data were 311 9° 212.3° 304.3° 221 ° 1.0 I 0 1 0 1.0 normalized to 1 by dividing by the maximum pressure amCase 34 3° 1 14 9° 24.9° 122.3° 1.0 1.0 1 0 1.0 plitude squared for each curve. Figure 2(a) shows the results III 302 3° 204.9° 294 9° 214.3* 1 0 1 0 1.0 1.0 .3° for the case with uniform excitation sources. There was good ■11 9° 124 3° 32 131.9° 1 0 1 0 1 0 1.0 agreement between the simulations and experimental results 158 0° 142 4° 59.2° 68.0° 1 1 I 2 1.0 1.1 for the main beam. When the array was excited by various Case 149 2° 133.2° 43 2° 52.4° l.o 1 0 1.0 1.2 phase settings, the simulations reasonably predicted the side- IV 232 4- 223 2° 313 2° 329.2° 1 2 1 0 1.0 1 0

248 0° 239.2° 322 4° 338.0° 1 1 1 0 1.2 1.1 lobe distributions found in the experiments [Figs. 2(b)-2(d)] The maximum power output measured was at least 14 for each transducer element.

B. Simulation results axis when the pressure at all of the control points were in phase [Figs. 3(a) and 3(b)]. Four focal spots were obtained if

The phased array ultrasound transducer used in the exthe pressure at the control points had 90.0° phase shifts from periments was simulated first. A calculation volume of 30 one point to another [Figs. 3(c) and 3(d)]. Comparing Figs. X 30X90 mm³ was used for this phased array. The axial dis3(c) and 3(d) with Figs. 3(e) and 3(f), the peak to peak distance from the center of the transducer to the water-tissue tance was increased when the control points on a 1.25-mm interface was 70 mm. The in^verse technique results are radius circle increased to a 3.25-mm radius circle. However, shown first. The amplitude and phase settings calculated by if the control points were on a 5.25-mm radius circle, the the inverse technique corresponding to selected control focal spots were at the same location as Figs. 3(c) and 3(d). points arc given in Table III (only the first four are shown) The field distributions (not shown here) were very similar to The contour plots of power deposition calculated for these Figs. 3(c) and 3(d). For ix control points, six focal spots array settings are displayed in Fig. 3. For four control points, were obtained at control locations with different patterns only one focus was produced, which was centered on the [Figs. 3(g) and 3(h)], and two additional focal spots were at the center. When the number of control points increased to eight, the power deposition pattern had the maximum pressure amplitude in froht of the focal plane with strong peaks at the water-tissue interface [Figs. 3(i) and 3(j)]- These results were used to select simplified amplitude and phase settings which were then used in necrosed tissue volume simulations (Table IV).

The isothermal doses for tissue necrosis for five amplitude and phase settings are displayed in Fig. 4. Let us consider the ultrasound pulse duration 10-s case (solid line). For the uniform excitation case, the focus was located at a depth of 59 mm from the interface. Based on the 240-miπ isothermal dose [Figs. 4(a) and 4(b)], the shape of the necrosed tissue volume was close to an ellipsoid. The calculated volume was about 85 mm³, measuring 3 mm laterally and 18 mm longitudinally. Isothermal doses are also shown m the same figures for the ultrasound pulse durations of 5 and 1 s. where

the input power was adjusted so that the maximum tempera¬

X-Distance (mm) ture was kept at 80 °C. For case I of Table IV, the necrosed tissue length [Fig. 4(c)] was 42 mm, and the width [Fig.

Fio. 2. Relative pressure amplitude squared distribution as π function of 4(d)] was 3.4 mm. The length was more than double and the lateral distance at the acoustic focus (; = 129 mm). The profile with ihe width was slightly enlarged compared to the uniform excitadotted line was measured, and the profile with the ιolid line »a> imulated The amplitudes were all the same vith phases (a) uniform. "*'. ⁽ and ⁽d) tion case. The computed necrose tissue volume was about various 339 mm³. Figures 4(c) and 4(f) snow case II of Table IV, for

Medical Physics, Vol. 22, No. 3, March 1995

FlG. 3- Contour plots of power deposition for calculated amplitude and phase settings. Th' ^■^nt^r•^, points used io calculate amplitudes and phases from Table I were (a), (b) case 1; (c). (d) case II; (e). (f) case III; (g), (h) case V; and (i), (j) case Vl. llie figures on the left side are axiaj plane distributions, and on the right side are focal plane distnbutions (z = 129 mm).

Medical Physics, Vol.22, No.3, March 1Θ95 302 X. Fan and K. Hynyn«...l Control of necrosed tissue volume during ultrasound ft. st 302

TABLE IV, The simplified phase and amplitude settings for the 16 square- To understand the effect of radius of curvature on the element spherically curved phased array used^* in the necrosed (issue volume necrosed tissue volume, the isothermal doses for case IV of calculations. Table IV, with an ultrasound pulse'duraiion of 10 s and op¬

A erating frequency of 1.4 MHz are displayed in Figs. 5(c) and

', ! 5(d). The distance from ihe transducer to the interface was

0.0° 0.0° 0.0° 0.0° uo 1.0 1.0 1.0 varied according to the radius of curvature to maintain a

Case 0.0° 0.0° 0.0^* 0.0° 1 .0 5.0 5.0 1.0

I 0.0° 0.0° 0.0° 0.0^: 1.0 5.0 5.0 1.0 constant focal depth. The necrosed tissue volume increased

0.0° 0.0' 0.0* 0.0° 1 .0 1.0 1.0 1.0 rapidly along the longitudinal direction toward the interface as the radius of curvature increased. However, the diameter

90.0° 90.0" 0.0° 0.0' 1.0 1 .0 1.0 1.0 of the necrosed tissue volume increased only slightly in the

Case 90.0* 90.0' 0.0° 0.0° 1.0 1.0 1.0 1.0

II 180.0' 180.0° 270.0° 270.0° 1.0 1.0 1.0 1.0 focal plane.

180.0° 180.0° 270.0° 270.0° 1.0 1.0 1.0 1.0 To show the capability of the phased array to move the focus, contour plots of the power deposition for the maxi¬

270.0° 180.0° 270.0° 180.0* 1.0 1.0 l.o 1.0

Case 0.0° 90.0' 0.0° 90.0° 1.0 1 .0 1.0 1.0 mum displacement are displayed in Fig. 6. The simulation

III 270.0° 180.0° 270.0° 180.0° 1.0 1.0 1.0 i .O model was the same as the one used to generate Fig. 3. The

0.0° 90.0= 0 0° 90.0° 1 .0 1.0 1.0 1.0 phase distributions were obtained based on direct calcula¬

270.0" 240.0° 210.0° tions, i.e., phase θ,, = 2 τtd„/λ, « = 1 ,2,...,^, where d_h is the

180.0° 1 .0 1.0 1.0 1.0

Case 300.0° 90.0° 0.0° 150.0° 1.0 1 .0 1.0 1.0 distance from the center of the element to a desired focus.

IV 330.0° 180.0° 270.0° 120.0° 1.0 1.0 1.0 1.0 Figure 6(a) shows the results for the case with uniform am¬

360.0° 30.0° 60.0° 90.0° 1.0 1 .0 I .O 1.0 plitude and phase setting. The focus could be moved 7 mm closer to the interface [Fig. 6(b)], or 7 mm deeper [Fig. 6(c)] compared with Fig. 6(a). Figure 6(d) shows the phased array which the necrosed tissue length and width were 20.8 and focus shifted 1.5 mm off the central axis. When larger dis5.1 rnm, respectively. The length was slightly longer and the placements were attempted, the phase increment between adwidth was almost double compared with the uniform excitajacent elements exceeded τr/2. Distributions generated by tion case. Notice that this case is equivalent to the situation those phases are no longer single, strongly focused ultrain which the phased array had only four elements with sides sound fields. Therefore these were the maximum displaceof 39.4 mm. The calculated necrosed tissue volume was ments achievable with this array in the sense of keeping the about 377 mm³. Figures 4(g) and 4(h) indicate that there fields in the region of focus similar to those for the uniform were four focal points in the power deposition for case III of excitation case. Table IV. The shortest peak to peak distance was about 4λ, Finally, Fig. 7 shows the input power requirements for the where X is the wavelength. The whole volume was the sumuniform excitation and four amplitude and phase settings mation of the four individual small necrosed tissue volumes given in Table IV. For ultrasound pulse durations of 5 or 10 which were produced by the four focal points. Each small s, the power ranged from 30 to 250 W. For a 1-s pulse, the volume was about the same size as the volume in the uniinput power could be as high as 900 W, depending on the form excitation case for the 1- and 5-s sonications. For the amplitude and phase settings. 10-s sonication a united volume was created with the total volume of 810 mm³. For the amplitude and phase setting of IV. DISCUSSION case IV in Table IV, the power deposition (not shown here) The experimental and simulation results showed that a indicated that there were four strong focal points (depth 54 16-element phased array can offer significant control over mm) surrounded by four small focal points. The shoπest disthe size and shape of the necrosed tissue volume during ultance between the strong peaks was about 2λ. Simultaneous trasound surgery. The experiments showed that such an array focusing at these points was performed to enlarge the heated can be constructed and it can deliver enough power for survolume compared to the uniform excitation case. The calcugical purposes. The simulated and experimental ultrasound lated isothermal dose [Figs. 4(i) and 4(])] from the temperafield distributions agreed reasonably well. The differences ture distribution had a maximum length of 38.5 mm and a may be due to uneven element size and power output in the maximum width of 7.2 mm, and the necrosed tissue volume phased array as well as measurement errors. Since the peak was close to 1.4 cm³. to peak distance was only a few millimeters, the ultrasound

To illustrate the effect of frequency on necrosed tissue detector may contribute some measurement errors. The simuvolume production, the isothermal doses for case IV of Table lation model accurately predicted the locations of sidelobes IV with an ultrasound pulse duration 10 s are shown in Figs. and main beams except that it over predicted the magnitude 5(a) and 5(b). The calculated volume was 60X60X90 mm³ of the sidelobes. When adjacent elements of a phased array for 0 5 MHz, 40X40X90 mm³ for 1.0 MHz. and 30X3OX9O are very close to each other, the mutual coupling between the mm³ for 1.4 and 2.0 MHz. The distance from the transducer elements would have an effect on the ultrasound field. The to the interface was kept the same (70 mm), while the fredifference between the sidelobes in the experiments and the quency was varied. The ratio of the necrosed tissue length to simulations may be due to the lack of such crosstalk in the width was almost independent of frequency. The size incalculations. creased as the freque_>._*.-y decreased. Tne shapes were similar The άυove results proved that the large element spuiri- for frequencies of 1 to 2 MHz. cally curved phased array can enlarge the necrosed tissue JUJ X- Kan ana K Hynynen- Control ot ru Isβa tissue volume during uitrasouno surgery JθJ

Fic 4 Isothermal doses for the uniform excitation case (a), (b) and the amplitude and phase settings in Table IV (c). <d) case 1. (e) (f) case II. (g). (h) ease IU: and (i). 0) "** ^V "T e figures on the left side are axial plane dismbutions and on the npht side are focal plane distπbutions [(b) : = 129 mm. Cd) 2= 1 28 mm. (0 z°' 1 2B mm. (h) z ^m ' 29 mm, and (j) ;= 124 mm] The ultrasound pulse durations used in the thermal dose calculauoπs were 10 s (solid hne) S s (clashed line), and I s (dotted line)

Medical Physics, Vol.22, No. 3, March 199S "« X. Fan and K. Hynynei lonlrol of necrosed tissue volume during ultrosαurtd.sur

Fio. S. The isothermal dose for case IV of Table IV «.nh various frequencies (a), (b), and with various radii of curvature (c). (d) The ultrasound pulse durauon used in the thermal dose calculations was 10 s The distance from the transducer to ihe tnierface was 130. 100, 78 and 32 mm for radii of curvature of 200, 160. 130, and SO nun respecnvely (a), (c) is axial plane distribution, and (b), (d) is the focal plane distnbution

volume. The phased array can also enlarge the necrosed tisThe 16-element phased array can generate four focal sue volume in only one direction at a time, if desired. It is points with a peak to peak distance as short as two waveimpoπant to be able to control the focal spot size so that a lengths. The maximum distance between the closest peaks is large tumor could be treated in a reasonable time. The conlimited to about four wavelengths. When the four control struction of the whole system was relatively simple due to points were on a circle of radius 5.25 mm, the phase increthe small number of elements. The 16-element phased array ment between adjacent elements exceeded π for this array. can also shift the focus off the central axis or move the focus Therefore, it is physically impossible to use the phase and along the central axis. However the distance is limited to amplitude setting obtained by inverse techniques to produce ±1.5 mm laterally and 14 mm in the axial direction. In movthe four focal spots at the control points. Although Eq. (4) ing the focus along the central axis, the array is similar to a represents an underdetermind system, it produced a solution concentric-ring array with two rings. Theoretically, the maxiat the control points. When muluple foci are separated by mum phase increment between adjacent elements is rt distances larger than 4.6 mm. the necrosed tissue volume Therefore, the maximum possible phase difference between becomes a few individual smaller volumes, which arc not the smallest and largest phases is it when moving the focus united for short ultrasound pulse durations along the central axis. From geometric considerations, this By decreasing the frequency, the necrosed tissue volume phase difference produces displacements along the central can be enlarged because the focal spot increases due to the axis of up to 23 mm (10 mm closer, 13 mm deeper). In increased wavelength. But the ratio of the necrosed tissue shifting the focus sideways, it is similar to a cylindrical- length to the width was kept almost the same. As the radius section array with four elements. The maximum possible of curvature is increased, the necrosed tissue volume can also phase difference between the smallest and largest phases is be enlarged. The necrosed tissue volume is increased mainly 3ιr for shifting,ιhe focus sideways. Geometrically, this phase in the axial direction. This agrees with previous expenence difference shifts the focus 3.4 mm laterally. However, the using single element spherically curved transducers. simulations showed that the phase increment between the Phased arrays require more input power than similar adjacent elements shoi'" * be less th»n ττ/2 to generate a "mgle focused traπsriiicrrs to reacu the same temperature single strongly focused ultrasound Meld level for the same ultrasonic pulse duration due to increased

Moitleol PhvsiCB. Vol. 22. No. 3. March 199S

Z-Distance (cm)

ACKNOWLEDGMENT

Fio. 6. Contour plots of power deposition for various amplitude and phase settings, (a) is the uniform excitation case. The

distance from the This study was supported by NCI Grant No. CA 46627. transducer io the focus was (a) 129 mm. (b) 122 mm. ( l 136 mm. and idl 129 mm. shifted 1 .5 mm

"'Department of Electrical and Computer Engineeπng. University of Arizona. Tu sun. AZ 8572) .

'K. Hynynen, "The threshold for thermally significant cavitation in dog's thigh muscle in vivo " Ultrasound Med. Biol. 17. 157- 169 (1991). focal spot size. The test array measurements showed that -J. P. Do-Huu and P. Hanemann. "Annular array transducer for deep therequired power levels can be generated in practice at least acousiic hypeπhermia," Ultrasonics Sy p. Proc. IEEE-81 CH 1689-9,

J - and 10-s exposures and thus, the proposed technique is 705-710 (1981). ible. ^?K. B. OcheJiree, P. J. Benkeser. . A. Fπizell. and C. A. Cain. "An ultrasonic phased array applicator for hypeπhermia." IEEE Trans. Soπics

Irt Order to obtain the desired field pattern, several field Ultrason.. SU-31. 526-531 ( 1984). control points are necessary to perform the inverse calcula. C. A. Cain and S. Umemura. "Concentπc-ring and sector-voπex phased tions. However, the number of utilizable field control points aιτay applicators for ultrasound hyperthermia." IEEE Trans. Microwave Theory Tech., MTT-34. 542-55 ) (I9S6). is limited by the number of phased array elements. Hence the ^SP. J. Benkeser, L. A. Fπzzell. K. B. Ocheltree. and C. A. Cain, "A tapered number of field patterns that produce significantly different phased array ultrasound transducer for. hyperthermia treatment." IEEE shapes or sizes of the measured volume is limited. Although Trans. Uluason. Ferroelec. Fr q. Conlr. UFFC-34. 446-453 (1987), only few amplitude and phase sellings were presented in this ^JE. S. Ebbini. S. Umemura. M. Ibbini, and C. A. Cain. "A cylindrical- section ultrasound phased-array applicator for hyperthe ua cancer study, the results provided information which could aid in therapy," IEEE Trans. Ultrason. Ferroelectr. Freq. Contr. 35. 561 -572 planning ultrasonic surgery using a phased array system with • (1988)^'. a small number of elements. These results have not been ⁷E. Ebbini and C. A. Cain. "A spherical-section ultrasound phased array optimized; nevertheless they illustrate the feasibility and applicator for deep localized hyperthermia," IEEE Trans. Biomed. Eng. BME-38. 634-643 0991 ). range of focal size and shape obtainable with a 16-element "E. Ebbini and C. A. Cain. "Multiple-focus ultrasound phased-array patarray. The calculations and experiments can be used to obtain tern synthesis: Optimal d ing-siπna) distributions for hypeπhermia." a set of different focal spots for a given array. This limited IEEE Trans. Ultrason. Ferroelectr. Freq. Contr. 36. 540-548 0989).

*M. Buchanan and K. Hynynen. "Design and experimental evaluau^'on of number of focal spots can then be utilized to optimally cover intracavitary ultrasound phased array systems for hyperthermia." IEEE the target volume in minimum time. The minimization of the Trans. Biomed. Eng. BME-41. 1 )78- 1187 (1994). duration of the treatment is an important factor for control^I0H. T. O'Neii. "Theory of focusing radiators." J. Acousl. Soc. Am. 21. ling the cost of the procedure, especially when MRI is used 516-526 (1949).

"X. Fan and . Hynynen. "The effect of wave reflection and refraction at io guide and monitor the treatment.¹⁶ If more control over soft tissue interfaces dunng ultrasound hypeπhermia treatments." J. the field pattern is required then an array with a larger numAcoust. Soc. Am. 91. 1727- 1736 0992). ber of elements is required. Such an array could al50 be used "X. Fart and K. Hynynen. "The effects of curved tissue laycπ on the

Medical Physics, Vol. 22, No. 3, March 1995 306 X. Fan and K. Hynynen: Control of necrosed 'laeue volume during ultrasound surgery 308 power deposition pattern_* of therapeutic ultrasounu beams," Med. Phys. rameters on high temperature ultrasound hypeπhermia. " Ultrasound Med. 21. 25-34 (1994). Biol. 16, 409-420 (1990).

¹³C. Damianou and K. Hynynen. "The effect of various physical param^liS. A-- Sapareto and W. C. Dewey, Thermal dose deierminaiion in cancer eters on the size and shape of necrosed tissue olume during ultrasound therapy.'- Int. J. Radiat. Oncol. Biol. Phys.-i0, 787-800 (1984). surgery..'- J. Acpust. Soc.- Am. 95, 1641-1649 (1994). "K. Hynynen, A , Darkazanli. E . Unger, and J. Schenck. "MRI-guided ^lJB. E. BiUard. K. Hynynen. and R. B. Roemer, -Effects of physical panoninvasive ultrasound stfrgery." Med. Phys. 20. 107-05 (1993).

Claims

CLAIMS :

1 1. An apparatus for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising

A. a plurality of transducers,

B. an excitation source for driving the plurality of transducers to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain,

C. the excitation source driving at least selected transducers at differing phases 1 with respect to one another.

3 2. An apparatus according to claim 1, wherein 4 5 A. the excitation source drives each of the selected transducers at a phase that 6 compensates for a phase shift effected by the skull in the ultrasound generated 7 and transmitted to the selected region by that transducer, 8 9 B. so that the ultrasound generated by the selected transducers arrive 0 substantially in phase with one another at the selected region. 1 2 3. An apparatus according to claim 2, wherein the excitation source drives each 3 of the selected transducers at a phase so that the ultrasound generated by the 4 selected transducers arrive at phases within 90° of one another at the selected 5 region. 6 7 4. An apparatus according to claim 2, wherein the excitation source drives each 8 of the selected transducers at a phase so that the ultrasound generated by the 9 selected transducers arrive at phases within 45° of one another at the selected 0 region.

5. An apparatus according to claim 2, wherein the excitation source drives each of the selected transducers at a phase so that the ultrasound generated by the selected transducers arrive at phases within 20° of one another at the selected region.

6. An apparatus according to claim 1, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.01 MHz to 10 MHz.

7. An apparatus according to claim 6, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.1 MHz to 2 MHz.

8. An apparatus according to claim 6, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region at a sonication duration ranging from 100 nanoseconds to 30 minutes.

9. An apparatus according to claim 8, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region with continuous wave operation.

10. An apparatus according to claim 9, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region with burst mode operation, where burst mode repetition varies from 0.01 Hz to 1 MHz.

11. An apparatus for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising

A. an ultrasound transducer, B. an excitation source for driving the transducer to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain.

12. An apparatus according to claim 11, wherein the excitation source drives the transducer to deliver ultrasound to the selected region at a frequency ranging from 20 kHz to 10 MHz.

13. An apparatus according to claim 6, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.02 MHz to 10 MHz.

14. An apparatus according to claim 13, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region at a sonication duration ranging from 100 nanoseconds to 30 minutes.

15. An apparatus according to claim 13, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region with continuous wave operation.

16. An apparatus according to claim 15, wherein the excitation source drives the plurality of transducers to deliver ultrasound to the selected region with burst mode operation, where burst mode repetition varies from 0.01 Hz to 1 MHz.

17. A method for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising the steps of

A. placing a plurality of transducers in the vicinity of an exterior surface of the patient's skull, B. driving the plurality of transducers to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain,

C. the driving step including driving at least selected transducers at differing phases with respect to one another.

18. A method according to claim 17, wherein step (C) includes driving each of the selected transducers at a phase that compensates for a phase shift effected by the skull in the ultrasound generated and transmitted to the selected region by that transducer, such that the ultrasound generated by the selected transducers arrive substantially in phase with one another at the selected region.

19. A method according to claim 18, wherein step (C) includes driving each of the selected transducers at a phase so that the ultrasound generated by the selected transducers arrive at phases within 90° of one another at the selected region.

20. A method according to claim 18, wherein step (C) includes driving each of the selected transducers at a phase so that the ultrasound generated by the selected transducers arrive at phases within 45° of one another at the selected region.

21. A method according to claim 18, wherein step (C) includes driving each of the selected transducers at a phase so that the ultrasound generated by the selected transducers arrive at phases within 20° of one another at the selected region.

22. A method according to claim 17, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.02 MHz to 10 MHz.

23. A method according to claim 17, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.1 MHz to 2 MHz.

24. A method according to claim 17, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region at a sonication duration ranging from 100 nanoseconds to 30 minutes.

25. A method according to claim 17, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region with continuous wave operation.

26. A method according to claim 17, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region with burst mode operation, where burst mode repetition varies from 0.01 Hz to 1 MHz.

27. A method for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising

A. placing an ultrasound transducer in a vicinity of an exterior surface of the patient's skull,

B. driving the transducer to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain.

28. A method claim 27, wherein step (B) includes driving the transducer to deliver ultrasound to the selected region at a frequency ranging from 0.02 MHz to 10 MHz.

29. A method according to claim 28, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region at a frequency ranging from 0.1 MHz to 2 MHz.

30. A method according to claim 29, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region at a sonication duration ranging from 100 nanoseconds to 30 minutes.

31. A method according to claim 30, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region with continuous wave operation.

32. A method according to claim 31, wherein step (B) includes driving the plurality of transducers to deliver ultrasound to the selected region with burst mode operation, where burst mode repetition varies from 0.01 Hz to 1 MHz.

33. An apparatus for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising

A. a plurality of transducers arranged in a two-dimensional array, and

B. an excitation source for driving the plurality of transducers to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain.

34. An apparatus according to claim 33, wherein the plurality of transducers are arranged in an array having any of a substantially circular and a substantially annular cross-section.

35. An apparatus according to claim 34, wherein the plurality of transducers are arranged in an array having a substantially circular cross-section.

36. An apparatus according to claim 34, wherein the plurality of transducers are at least one of (i) mounted in, and (ii) separated from one another by, a damping agent.

37. An apparatus according to claim 36, wherein the damping agent is any of a natural or synthetic rubber.

38. An apparatus according to claim 34, wherein the excitation source drives at least selected transducers at differing phases with respect to one another.

39. An apparatus according to claim 1, wherein excitation source drives each of the selected transducers at a phase that compensates for a phase shift effected by the skull in the ultrasound generated and transmitted to the selected region by that transducer, so that the ultrasound generated by the selected transducers arrive substantially in phase with one another at the selected region.

40. A method for delivering ultrasound through the skull to the brain of a patient, the apparatus comprising the steps of

A. placing a plurality of transducers in the vicinity of an exterior surface of the patient's skull,

B. driving each of at least selected transducers generate and transmit ultrasound through the skull, and determining a phase shift effected by the skull in ultrasound generated and transmitted by each such transducer, C. driving the plurality of transducers, together, to generate and transmit ultrasound through the skull to induce cavitation at least at a selected region of the brain,

D. the driving step including driving at least selected the transducers at phases determined in accord with step (B) so that the ultrasound generated by at least the selected transducers arrive substantially in phase with one another at the selected region.

42. A method according to claim 40, wherein the selected region ranges from 1 mm³ - 1 cm³ in volume.