WO1998007373A1 - 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 5A and 5B illustrate the effect of applying ultrasound in accordance with the invention to brain tissue.

Figures 6A 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), a copy of which is filed as an appendix hereto and 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 a a₂, a₃, etc., as shown in the drawing. These phase shift factors α,, α₂, α₃, 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,, a^ 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. 1 178-1187, a copy of which is filed as an appendix hereto and the teachings of which are incorporated herein by reference. Array steering, or focusing, is particularly discussed in that article, for example, at pages 1 179-1181 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 (MR ).

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 α₂, α₃, etc., having first and second components as described herein and above.

Referring to Figure 1 1, 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 PV pv [2] p is the density (ρ_b = 1030 kg/m³ , p, = 1380-1810 kg/m^J) (s= skull, b = brain), c is the specific heat of the medium (c_t, =3.9 kJ/kg/°C ; c_s =2.1-2.7 kJ/kg/°C), v is the speed of sound, and α is the amplitude absorption coefficient of the tissue (α_b= 5Np/m/MHz, α, = 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_bVP,²) = (α/α_bXpJp.)'(vJv.)(c/c.) = 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-^2P& = 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 Wcm'² and 2000 Wcm ²at 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 2431L ). 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 4 A, 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-

Frequencv (MHz) Amplitude fMPa f Pa (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, rhe 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. Ens., v. 41, pp. 1178 - 1 187 (1994)

Design and Experimental Evaluation of an Intracavitary Ultrasound Phased

Array System for Hyperthermia

Mark T Buchanan and Kullervo Hynynen

Abstract — For evaluating the feasibilit^y of treating prostate nations the depth 10 which thev can effectively heat, and cancer, a 64-element linear ultrasound phased arrai applicator their limited ability to control the power deposition helds for intracavitary hyperthermia was designed and constructed [2] [ 1 Both deeper penetration and increased control over Λ 64-channel ultrasound driving system including amplifiers. the powei deposition pattern can he achieved using linear phase shifters, and RF ower meters was also developed to drive the array. The design of the array and driving equipment are phased arr.ivs Since phased arrays CJΠ focu iheir radiated presented, as are the results of acoustical held measurements energy, they theoretically could heat tissues 10 therapeutic and in vino perfused phantom studies performed with the array. tempeiaturcs deeper lhan nonfocused arrays, The ultrasound Several techniques for heating realistically sized tumor volumes power dcpo-iuon pattern can be clecti omcalh tailored as _•vere also investigated, including single focus scanning and two techniques for producing multiple stationarv foci. The results ncce.ssarv to produce therapeutic temperatures within a desired show that the operation of the array correlated closely with v olume In .II CJS w ere heating would b undesirable, ihe the theoretical model When producing a single stationarv focus, phased ,ιrr_ could take adv . t.i c ol destructiv e interference the array was able to increase tissue temperature by I2°C in to miπiπii/.e power deposition at those locations. The ease vitro in perfused phantom With some minor improvements in bv which Ih υw'ci deposition pjttern can be electronically array design, intracavitary phased arravs could be o aluated in a clinical environment. .tlieicd in red time provides a means lor the compensating tor varying physiological parameter paiieni positioning, and fσi minimizing patient discomfort

1 INTRODUCTION Several types ol ulti asound phased arravs have been

INTRACAVITARY ultrasound arrays offer an attractive proposed or built I01 hypeπhermic purposes These include means of inducing local hyperthermia in deep-seated tu Unierπur and Cain^'s sccior-vouex and concentric ring applirnors located near body cav ities By locating Ihe radiators as cators |4], |5) the cylindrical section applicator developed

close as possible to the ireaimen; site, problems frequently Ebbini ct al 16). |7j. s well as the upcred arras developed encountered with external techniques, such as blockage of by Benke.scr ei al [8 | Each of these arrays is composed of the acoustic w indow b> bone or gas, or simply the inability anyw here from 16 10 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, nonfocused arravs of half-cylinder show significant potential, they ate meant to be used in external transducers operating at 1 6 MHz suggest lhal such arrays can applications and therefore aie unsuitable for intracavitary use be clinically useful in the treatment of prostate cancer ( I ] in their reported configurations

The proximity of the prostate to the rectum w ll makes it Previously, a thcoi eiical siudv on the feasibility of iπtracav- a good candidate for heating using intracavitary ultrasound iiarv phased arrays using half-cylinder elements had been done radiators. Since the prostate is located very near ihe anus and by Diedeπch and Hynynen [31 Based on many acoustical and only millimeters a*av rrom the rectum wall, an applicator can thermal simulations. 11 was concluded that a practical array be easily located close to the prostate The prostate is one ol would be composed of 30 half-cylinder elements wit 2.5 the most easily accessible tumor sues, 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 lhai the crating lobes formed with this useful. As such, most of the experiments were conducted u array could be negated by using surface cooling. One of the goal of heating the prostate in mind the major considerations in specifying ihis design was noi

While the nonfocused arravs have show n 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

M-miscπpi received November I " 109 rev isej Au_just V IW4 Thi_* rπιnιmi7^β the number of amplifiers led to the specification ol work was supported bv grjnt number O I CA 4S9'° Irorπ ihe Nmional a 30-elemem array Cancer Institute Initial array designs were based on some recommendations

The authors are <«'ith the Department of Rjdiolσ_jv DIVI_{MCΠ 01} MRI Bπgham and Women s Hoipnal H-π ard Medical School Boston M * 021 15 from the theoretical study and expanded upon in an attempt USA to construct a practical intracavitary phased array It became

IEEE Log Number 9406164 apparent that more than 30 amplifiers would be necessary for BUCHΛN»Λ AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

Fiv I The 64-clemeπι and a single element sho me the Rβomeirv ct the ,ιrr„. ThL- π.¹ aι mourned ro a brass shell 19 rrm in dumeier Thf a k of ihe applic-ior * a* removable so thai » ιrcs could be conπecied to me inn r cicum-c

a useful array This led to ihe development of the 64-cnanncl of conimerc ull*.

au_bιe eauipmeni lo απ e ihe arτa> 3 driving system to overcome the limitations imposed bv the 64-chanr.ci ampiiner sysicm <* as developed This system is number ot amplifiers This allowed for ihe evolution of a 6~- composed of phase

cscie controllers, amplifiers. channel phased arτ_v with 1 .73 mm cenιcr-to-ceπtcr spacing R F ow er mcicrs. rc impedance matching networks I see and a toiai arra>. length of 1 10.5 mm operating at 500 kHz bloc), diaruir. in F ι : 2 ι ana is controlled DV an IBM PC compatib le comnuie: . .a <ι rjigii.il I O interlace

11 M ATERIALS \ND METHODS Tne amDiiners arc b.ised on a swucning MOSFET αesign ana are ae' icncd arc.ιr.3 intcmationa. Rcciilier s IR21 ) 0 ua;

A λvt us Consn iic riπn MOSFET αrn er £ r. c ihe amplifiers is capable oi deli v ¬

A 64-elemenι array a a* constructed using half-cvlinαεr ering up to I W of RF po» er at 500 kHz into a 50-W load . nsducers operating in iheir resonant radial mode at 500 kH_ from each cnanncl cr _ :otai output power of about 850 W The array was made by slicing washer-shaped elements with The amplifiers

er: digital loϊic inpui signals into high a diamond wire saw (Laser Technologies. Noπh Hollywood power sine n a es « nue preserv ing the pnase of the inpui CA) from 15 mm O.D bv 30-mm long cylinders of PZT- ιenπl Th .inipntuαe of tne OulDUl signal is controlled by 4 material (EDO. Sail Lake City L'T) w ith silver eleciroαes ιw.o method-. All of tr.e amplifier outputs can be controlled plated on both the inner and outer w all suπaces. The transducer simultaneously ov aαiusunε tne output voltage on a 1000 W slices ere elueα togeiner using a silicone adhesive i Dow DC supply cr tne ou'O'-i ol each cnannel can DΪ ind dualh Coming. Midland. MI ) with 0.17-mm thick silicon rubber conirolled bv varying '.he duly v ele of tne input signal. The spacers !SPC Technology . Chicago. ILϊ beiwecn each elemeni rcsull oi reducing Ihe αutv evele of the input signal t the duly (producing a dead space of 0.23 mm oetween elements'). T e c\ cιe is tne percent of ' on^" time of the input signal per clock stack of elements as then cut in half along the axis o the ocie i is a correspor.a'.ne decrease in tne amplitude of tne cy nαe: and the two half-cylinder sections glued together to ouiDut signal form tne full array The array was bonded to a brass shell Since ihe ampiiner; require αigitai input signals, the phase lo form the complete applicator, as shown in Fig. 1 Wires shilling jn duiy-cvcie control is implemented using dizital were soldered to the inside wall of each array elemeni that .ouniers | i 0|. I 1 11. These circuits provide 22.5⁼ pnase shift extended the length of the shell to tne handle where they were resolunon Horn 0-360" . and l b steps of duty cycle (amplitude) each connected lo a 2-m RG-178 coaxial cable Tne electrodes ihat allows tne outpui to v ary from 0- 100% of the output on the ouier surface of the array were electrically connected by power, as allowed bv tne voltage from the DC supply While single wires embedded in beads of silver epox> iChomeπcs. ihe onasc shut resoiur.o.n mav seem somewhat limited, studies obum. MAi running alone bout edges of the array. n_2>e snown that n is sufficient for the uses described heir

1 1 1 1. ( 12 )

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

One of the primary disadvantages of phased arrays is Ihe by power meters mat measure both the forward and reflected increased complexity of the driving equipment. Due to a _lack RF power 1 131 The power meters, wnich were also designed /07373

.^

F'« Block diagram of ihe o.^.s«c *^^ - t h > MKipir ... ii fun*? amplifier

TΛ B LL I

and constructed in house, allov. for v erification of the output characterized using a hydrσpnσne and the results used to power on each of tne amplifier channels, and allow for compensate the excitation signals the monitoring of (he output; ^o that possibly dangerous fault conditions do not occur B^ monitoring the amount of reflected power, changes in the electrical matching and the C Ai. w:liC Moάclinv array operation can oc easily r- πitorsd so mat ej arrav The 3-D uhrasouno modeling program used * as ortginalh performance can be maintained developed b> Diedeπch and Hynynen (3j and is onlv bπerl)

Finally, to maximize pov. eτ im^ier betw een the amplifiers introduced here The rouline models Λ^' cylindrical radiators of and arrav elements, a simple LC :~,ι-.atτ zt matching network finite length, radius and separation as each of the ultrasound was used on each channel peiw÷ir. ;.-.= rx>« er meters and the transducers in the array, and assumes a target media n3vιnj array. This not onlv maximize: ;-v - er transfer, out ensures similar acoustical propcπies as ussue (Table I]. The routine greater umiormity in tne acous::;;. '. -TJI signals for a BIN en models the surface ot each of the cylindrical elements as an excitation p.naje ana amplitude i.-;.:c r^peπmeπtally b> a evenι> spaced grid of simple hemispherical sources and uses needle hydrophone ■ \λ'ιtho-ι tn .-~. zi. matcnin . tne pnase Huvgeπs principle to sum tne contributions of each source at shift and amplituαe distortion i- .-t; r e impedance decacn poini in the held All of the simulations ere αone 'Uh ferences bei i cen lements :n '-'.τ -.-; • « ould have to be Ihe 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 ayleigh-Sommerfeld integral for using an arrav wnh Λ^' elements, and a field of M control pom Pi')

Pi z. r . ^I Σ Σ ^^< ( 1 ) jp k -/H'...

.: ι=ι = — ^u" ∑ - ^■dS (4) where where t> is the densuy of the propagation medium, c is the

P. I 7 ) = P.. ,j(2'/ - 2_* ÷ - *) speed ol sound in the medium, k is the wavenumber. 5' is ihe

(2) surface of the source, u-, is the particle velocity normal to ihr surface of ihe source, and r_m- r„ is the distance beiween . and n. is the number of sources in the

lit is irte control poini ; r-, ι and the surface cf a source (r„ ). - 1 . number of sources in the theta direction. P. is the pressure

Z W . and r = 1. Z Λ^'. from a singie source (Pa), P,_a is me pressure amplitude at the

To simplih programming of this algorithm, (4) can be surface of the source (Pa). r„ is ihe radius of the source <m), r described b> is the radial distance from the center of the source (m), / is the wavelength (m). 6 is the phase of the excitation signal Irad), Hu = p (5 ) / is the operating frequency (Hz), and α is the attenuation coefficient (Np/m) where u is ihe excitation vector, p is the complex pressur. M ihe conirol points, and H is (he remainder of (4). For intracaviiary nypeπhermia uses. Ihe number of elements in

D Focusing Techniques the array i ts") is always greater than the number of control

The single focus case is the simplest Jorm ot to using that points ( M) This leads to an undcrdeienmtned system of can be done "* nh a phased array. The single focus is produced equations w ith an infinite number of solutions. The mu mum by senine ihe phases of the driving signals so thai constructive norm soiuiion I u i 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 stgnals for an array with ,Y elements can be calculated from ihe differences in tne path lengths d. between = H" (HH' (6 ) each array element and the focal position bv wnere II" is ihe complex conjugate of H With this technique. all of ihe arrav elements contribute 10 all of the foci, unlike

Δ-i,

2- ihe _SΓ_JI _II focusirir lechntque where each element contributes to only one focus where , is tne phase tradiansl of the ^{• ' h} element. / i.s the wavelength (mj. ( = 1 , 2 ;V and x;ι = 0. 1, 2. . . . The size 111. EXPERIMENTAL TECHNIQUES of the focus produced with this technique is usually too small to heat an entire tumor volume, and therefore other techniques A Λt.υusιu al Efficiency Measurements must be used to heat larger volumes.

The acoustical output of the cylindrical transducer elements

To heal larger volumes, two techniques were investigated, was measured using a radiation torce technique [ 16J. The single-focus scanning and mulnple focusing. The single focus arrav w as placed in a brass cone with 45^c Sides to reflect scanning routine simpiy stepped a single focus back and forth ihe radialiv emitted ulirasound fteids down into ihe acoustical along a predetermined length of the arrav The focal depth was absorber [ 17). Tne force on ihe aDsorber was measured using kepi constant while the focus was stepped every 300 ms to the a Metier AE 160 (Hightstown NJ l 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% ol maximum power power meter and Ώ Werlatσne C2625 (Brewstcr. NY⁾ dual at the center of the scan to flatten the temperature distribution direcnonal couoiei The efficiency was calculated as the ratio in the perfused phantom expeπments (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 loci were calculated The ultrasound fields were mapped in a lank of degassed^, using two techniques, solit focusing and the pseudo-inverse. deionized waier by mechanically scanning a thermocouple To create mulnple foci with the split focusing technique, the embedded in a small (2 mm diameter) plastic sphere The ther^¬ array was divided into subarrays. eacn of whicn produce a mocouple was positioned bv a three-axis computer controlled single focus in ihe same manner as previously described. The scanning table The applicator was mounted on a rotational pseudo- inverse method, developed by Ebbini and Cain ( 15). device that allowed measurements to be made in a radial arc uses a series of control points that represent the magnitude around ihe array by rotating the array. Measurements were of the ultrasound field at given points. A brief summary of made on a 1 x 1 mm gnd with the recorded data being the Ebbini and Cain s tecπnique follows. average of three consecutive measurements. I It- IEEE TRANSACTIONS OS BI0MED1CΛL.E«G, ING. VOL -1.-A.0~i:. DECEMBER. UWJ

C. In Vttro

Experiments

Alcohol fixed canine kidneys were used as phantoms far studying the heating characteristics of the array. The kidneys had previously been prepared as described by Holmes et al. [IS], and were rehydraied pnor 10 use. The experiments were conducted at room temperature using degassed, dcionized water as the perfusate A metering pump (Fluid-Meteπn Inc RH1C C. 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 bv gently sandwiching it between two PVC membranes mounted to a Plexiglas frame. The applicator was πrmlv clamped to the

frame to maintain a fixed distance between tne surfaces of the Fip 1 Du? ram o⁽ ilic ... , kidney t»penmenul setup kidnev ana array Fig 3 shows diagram of the experimental setup

Temperatures were measured in the kiαnev ubing either seven sensor probes sutured in place perpendicular (o the array through the focal region, or by one or two single sensor thermocouples pulled along a path parallel to tne 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 arrav as 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 noi

veiv nne spatial detail of the temperature distributions the pull-back technique was more frequently used

The pull-back experiments were conducted b\ pulling one or two single uncoated thermocouples (0.05-mm »M re J bv. .ι computer controlled stepper motor along a track parallel to t 'ici. ic si Cow*' | W | the arrav tnrough the kidney nssue The temperatures were p -- l A. ou tc ai ou'Dui α^ci *s i f unction ol eteciπcj" mcji power to^r measured even, I mm and the average ot tnree rejomss *a< a ' iπclc ju mm ^• , i c \ ι no;* element ana _rτ_vs rrnae ur υl ! ' _ . mm jn° ^ m ei errrnn .i r j u . mm inMjiatni B tw een cicmenr recorded Prior lo each of the experiments a oaieitne ιem perature was established along the path of the inermocouplci The kidney was exposed to u rasound for 20 min 10 allow i width < >~dveιeπsιh ι u hich would proouce

spherical the temperatures lo reach steady-state before the temperature wd elror. i ~I h; colnmation is necessarv to assure tne ai ee profiles were measured. The difference between ihe baseline reflected tn inc reflecior are normallv incident to tne surtace and final measurement w_s used to calculate the lemperaiυre of ihe aosorcer rise The kidnev was allowed to cool to room temperature 1 he lesuits or inese experiments are sho_*n in Fie None (typically . 30 mm) before to the next experiment wa started of (he arravs eλhibneα verv nigh etncieπcies despite the lou The mermocoupics were mainh located in the medulla or operating

Tnough not shown here some resul the kidnev The steady-state temperatures in the medulla have were verified us e caioπmeiπc lecnniques While tne 30-mm been shown to be a strong function of the flow into the kidnev long full cylinder element had an etficiencv of 71 ty the arrav s [19]. The flow values were kept relativelv low in order to enmbited efficiencies o^f 27% and 17% for the 2 2 mm and 1 5 simulate the perfusion in the prostate 120], [21 ] mm arrav s re pective These ef ficiencies indicate mat acout 80^ of the eleciπcai Dower delivered to tne I 5 mm wide

IV RESULTS elements u^red in tne pnaseα arravs is lost either as heai in the transducer or in the electrical matching and transmission lines between tne a plinc and ihe array

A Acoustical Effluents Measurements

The acoustical efficiency of the 500-kHz PZT--1 transducer

B L Hi as una Field Meaniicments material was measured for Λ full cylinder 0 mm Ions and for two half-cyliπacr arravs wilh 2 5 mm ana I % mm center The eitccts of element spacing were stuαied DY measuring to-center element spacing ι2.2 mm and I 5 mm elemen^; the ultrasound held distrirjutioπs produced ov two io-eιemen' widths) w iiπ total array lengths of 4 7 mm and 2S 5 mrr, arrays VMIΓ. 1 h mm and 2 5-tnm center-ro-cen-c- e'ement respectiveK Small arrays were tested so that ;nc resulting spacing F- ic 3ι a i and l O i snou- s the acoustical field plots made field would De collimatcd unlike a single arrnv eiemeni in a plane parallel to tne arras and normal to tne surtace ol BUCH»S »Λ AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM 111}

Aa<j«i Oman-it ι m> HaO M otR-cm (mm)

( J jr.J I 5-mm i r> • ctemcm center ιo-cenιcr ip-απ? _nd iπeir correspondmc iimυuuor.

nd 1 1 mm trom me ceπirαi ins

the hall cylinder elements for two 16-eletnent array with 1 S- So far. all of the acoustical field measurements presented mm and 2.5-mm element spacing along w ith their respective have been in a plane parallel to the length of the array and simulateα fields (Fig. 5(c) and (d)). Both arrays were focused normal to [he surtace ot the arrav Ideally, the acoustical field 30 mm from the array surface and 15 mm above the central would taper on toward Ihe edges ot tne array but be more or axis of the arrav. Tne most significant etlect illustrated here is less uniform -round the arc of the array Unfortunately , this the generation of the grating lobe bv the 16-element array w ith ideal does not match reality, as cart be seen in Fig 7 The 2.5-mm center-to-center element spacing (Fig. 5(b) a d (di) 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 tne position of the thermocouple ιr, of the locus but also could cause heating in unwanted areas, the foc.il reuion and rotating the arrav about its aλis. The arrays must be designed to minimize grating lobe formation. acoustical intensity is not at all uniform ana. in fact, vanes The array with l.S-mm center-to-center spacing produced as much as 50% before tapering off at the eoges. All of the much better field distributions than (he array w ith 2.5-mm 2-D acoustical field plots shown nere were made at the 0' cemer-to-ctnter spacing. Therefore, in order to ha e center- rotation anele where tne intensity is only about 50% of the to-ce er spacing 1 S 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 * as reduced to 0.23 mm alters the peak intensity and does not effect tne overall snapc

The complete measured and simulated plots of the ultraot tne acoustic field sound field produced by the 64-elemenι array with 1.73-mm Fig δtai and (b) snows the acoustical field plots produced element spacing are shown in Fig. 6(a) and (b). respective!) usine the two multiple focus techniques: pseudo-inverse and The arrav as focused 40 mm deep, and 30 mm from the spin locusinε Below them, in Fig. S(c) and (d). are theu 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 svnthesize foci 40 mm from the array suπace and 10 mm from its central axis. In fact, this array is capable o focusine to on either side of the central axis of the arra^y (20 mm apart) about 35^ς from the central axis of the array without prooucine using the 64-element array Notice that the pseudo- inverse significant grating lobes technique produces sharper foci than those produced by the I IM IEEE TRANSACTIONS ON BlO EϋlCl C trlG.t. JtlNtJ.'Vp 41. INO. 12. DECEMBER U*;.

AadMl Dc-Ucnc* (mm) R«Oc«l otnanca (mm) ibl

Fig 6 Field plot of the cV-eleme -nay ( l alone »tth Ihe iimulition results 'b> it locm *J mm Irom the vjrlaee of ihe »rr>y ind -_^■ mm off its cental «ΛIS

1.0 — the circulating water in the v>ater bath All of the following 0.9 — results were obtained in the center of the kidney >*here the focus WJS located and the surface effects were smallest. The o β - temperature rises achieved in the middle of the kidney indicate c ? - thar therapeutic temperatures can be achieved with a stationary a t — single focus at realistic perfusion levels Typical cross-section profile_*, measured in the middle ot^" the kidney with ine pull-

0 5 bcitK lecnmque tor . center focus and 15 mm on either side

0 4 — are shown in Fig. 9(b)

The tnree different techniques for heating larger v olumes

0 J - p>eudo-ιnverse. split focusing, and scanned single foci arc 0 2 compared in Fie 10(aι and (b) The t o multiple tocusine

0 1 technique w ere used to produce foci 40 mm from the surface of the _rrav and 10 mm on either side of the arrav s central 0 0 _d m;, while the cannine routine produced a series of Single

^■90 -7t -60 -4S .30 -IS 0 IS SO 44 «0 7S SO foci 0 mm trom the array surface wun a toiai span ol RoUtion.i Angle (Dnjr*κ) 20 mm The mermocouples were located 8 mm and 15 mm from the surtace of the kidney nearest to the arra>

Fie 7. Smcle dimensional radial field plot made »nhιn the focus -s . function of ιh« .ntie o( rouuυπ of ihe arr.) Tl-.e center σl the jrτ»c .r (55 mm and 42 mm from the surface of the array). The is defined as ihe lero point. scanning technique produced a narrower profile than the two multiple focusing techniques and the distribution Jacks the temperature drop between foci The two multiple focusing split focusing method This, occurs because the pseudo-inverse 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 snarper foci than ihe split focusing technique w hich, temperature rise on the deeper thermocouple (bevond the since it divides the array into t o subarravs, effectively uses locus i 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 lecnmque. z previously explained.

C In Vitro Kidney Experiments

Fig. 9(a) shows the temperature rice ver^u* time along a v DISCUSSION AND SUMMARY fixed seven-sensor thermocouple probe located perpendicular to the array The distances marked denote the distance of An intracav itary ultrasound phased arrav composed of half- the thermocouple from the edge of the kidney nearer to cylinder transducer elements has been constructed for inducing the array. Tne array, wit 32 active elements, was focused hvoertnermia in the prostate via the rectum. A 64-channei along us eentrai axis and 10 mm into the kidney . (Thιrt> - amplifier system has also been designed and constructed ic two active elements were used since the array length in this αm e tne phaseα

. As was shown by the acoustical field conήeuration could adeαuatelv cover the kidnev ; 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 ; ^{_ i} Note that the kidnev fields containing more tnan one focus The arrav is curremi'. is only 35 mm thick and thus, the temperatures close to the capable o! producing a I 2°C temperature πse in a oerfused surfaces of the kidney are dominated by cooling caused by phantom using a stationary focus, and smaller temoerature BUCHANAN AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM

FUadh-t ouunαe (mm) nadM omane-i |mm)

Fig 8 Acoustical field plots produced mine inc pseudo-inverse and tocusmg (b) tecnπioues and the related simulated rr<un< ιι.-ι an-i (<j)) T * loci were produced -0 mm deep, and 10 mm on either siαc oi tne ceπtial ans ot trie arr v

increases were achieved in larger volumes by scanning the experiment because the impedance of the array elements did focus or by creating mulnple stationary foci. This was achieved nor return to their original 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 arτa> included arcing oetween elements and water seepage ability to heat tissues to therapeutic temperatures in a regubehind the arrav mat reouced the array efficiency in subsequent lar and predictable manner. The peak temperatures achieved experiments until the array was repaired ana remaiched. With within the perfused phantoms tended to vary considerably from a more carefully cons ucted 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 in good agreement u. ιιh the perature profiles that compare various focusing techniques theory, and the teenmques for heaπng larger volumes were were normalized to better illustrate the overall differences all functional, though tne multiple focusing techniques were in the shape of the temperaiure 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 hyoerthermia system. thermocouples within the kidneys, the inefficiency o the arra> . 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 aιτay would usually drop radiators still lacks control in the angular direction t around the considerably during the first experiment, and during each subarc oi the array). Additionally, since the cylindrical radiators sequent experiment due to changes in the electrical lmoedance αo not have uniform angular intensities, the angular heating of the array elements T o primary factors were responsible pattern is somewhat degraded, though thermal conducuon will for the observed changes in the electrical impedance of the probably smσotn 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 duπne sonication. The for other cylindrical transducers [2], [ I I ]. Radial control second impedance shift that occurred from experiment to could be acmeved DV dividing the half-cylinder elements into *6 IEEE TRANSACTIONS ON BIOMΪOlCAl ENCIv^»»glNG. VOL 41 NO II. DECEMBER 1994

Anal Dlft-ftcct (mm)

Aiiβl Distance* fmnij

0 1C 30 30 40 l b

Ailtl Distance |fnm| Fip 10 Norm l i z e icmperaturc/pow er profile*; measurefl r -» perfused <h) kidnev tor mump c tocu- mp iump ine pseuαo-i erse .no focusing _technique. «c c-n j< msle-focus

The mulnple tr>c ι * :ιc created

Fig 9 Temocraiure profile measured in - erfused kidnev min ; SI activ e ιυ mm αeep ! <■ <-~ on either side ol ihe central a»ιs hπe "e scanning elements locused 30 mm rrom ihe arτ-v >,urfacc » hilc ihe - idnev » jc, perfused s done c i i .inp ir. tgcυ< 40 mm deep -nd scanned l u mm on eⁱther •i the rate o[ 2.9 kj rr,- ' ',- ' tj mL min " fl » ) The array wo positioned side of inc central a cu Tne me_«.urcmcnι_. were made & mm f a i and I - mm I mrn trom ihe kidney surtace The thermocouple location indicate the (b l from tne cun ac_t oi ihe kidnev w ith ihe arrav positioned 2 ' rr.m trom the distance from the surface of the kianev u'ι The astal distribution measured kidnev su'lscc 12 mm from the surface of Ihe kiopey uc g a pull-back, thermocouple (bl in a separate e-penmeni minimized grating lobe formation, higher frequencies would pie-shaped subelemeπts and driving each subelement indepenincrease the oo c er aosorption in tissues Since most tumor dently. However, this would not only dramatically increase sites lor wnich this applicator would be used occur near the number of amplifiers, but the resulting subeiemems would _th_ a_Mi v, ill the αeep penetration aliowed by the 500- have virtually no inner electrode due to the wall thickness kHz operating trequency is not πecessan, The difficulty in of the element This *ouid seriousiv decrade the already low increasing (he operating trequencv is that the elemeni size and acoustical efficiency of the cylindrical radiators center-io-center element spacing would neeα to oecrease as

A possibly better arrav design <* ould utilize planer array the frequenc increased making array design more difficult elements mounted or, a rotating platrorm By tilling tne array A* a conc_lusion tne intracavuarv. electrically focused array back and forth, ine same voiumc could be treated as wn did αemonstrafc a; practical depths tor prostate treatments the cylindrical radiators By controlling the acoustic field and (about i r r, ) the feasibility of using phased arrays for output power a; a function of tilt angle, 2-D control over intraca uan nypeπnermia purposes With a more carefully the power deposition field can be achieved wunoui requiring construc_ted

most of the power limiting problems ex additional ampliners An additional improvement on tne arrav perienced can be avoided The acoustical held Plots were design would be to increase the operating trequencv While tne in good agreement wn the tneorv. ana the tecr.niques for 500-kHz operating trequencv allowed larger elemeni sizes and πeaung larger volumes were all functional, thougn in: multiple BUCHANAN AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM 1117

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 uurasound surgery using a 16-element phased array

Xiaobmg Fan*¹ and Kullervo Hynynen

Department of Radiolory Bngham and V>amen t Hnspitul. Harvard Medical School Boston Massachusetts 02115

(Received 24 March 1994; accepted for pubL'cauon 31 October 1994J

Focused high-power ultrasound beams are well suited for noninvasive local destruction of deep target volumes In oroer to ovoid cavitation and to utilize only thermal tissue damage, .high frequencies ( 1 -5 MHzl are used in ultrasonic surgery However the focal spots generated by sharpl₎ focused transducers become so small that only small tumors can be treated in a reasonable lime. Phased array ultrasound transducers can be employed to electronically scan a focal spot or to produce multiple foci in the desired region to increase the treated volume. In this article, theoretical and experimental studies of spherically curved square-element phased array s for use in ultrasonic surgery were performed The simulation results were compared witn experimental results from a 16-etement 2jτay It was shown that tne phased array could control the necrosed tissue volume

using ciosely spaceα multiple foci Tne phased array can also be used to enlarge a necrosed tissue volume in only one direction at a time. i c . lateral or longitudinal. The spneπcally curved 16 square-eiement phased array can produce useful results by varying the phase and amplitude setting Four focal points can oe easily generated with a distance of tw o or four wav elengths btiween the two closest peaks The maximum necrosed tissue volume generated by the arrav can be up to sixteen times the volume induced by a similar spherical transducer Therefore the treatment time could be reduced compared w ith single transducer treatment.

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

In ulirasound surgery, to avoid cavitation and have a sharp and electronic circuits are also required boundary between the rumor and normal tissue, high freThe motiv ation of this article was to demonstrate that a quency focused transducers have to be used ' The focal spots phased array system ith a small number of elements can be generated bv snarply tocused ultrasound transducers are utilised lo provide control over the focal spot size for surgismall when compared with the diameter of manv tumors cal purposes Tne excitation signals of tne elements vary in Obviously this tvpe of focus is not efficient to treat large both amplitude and pnase Hence the aior tasic in utilizing tumors, which would require a large nu Der of exposures to phased arrays is to determine tne phases and amplitudes cover the whole target volume. However, in order to accuneeded to produce desired ultrasound fields. The inverse racy treat the target volume close to critical structures, technique introduced bv Ebbini and Cain⁶ can be used for s . focal spots may be required duπng pa of the therapy direct synthesis of ultrasound fields with multiple foci. Sev Thus, controllable focal spot si2e is required. The most ateral amplitude and pnase 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 ihe numuse phased arrays ber of utilizable control points was also investigated for a

Phased arra applicators were introduced to ultrasound given array Several simplified amplitude and phase setting? hyperthermia cancer therapy in the early 19S0's. During the based on the calculated amplitudes and phases wxre empast decade, many efforts have been made to inv estigate the ployed for ultrasound fielc calculations Tnesε dπving signal advantages of phased arravs 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 nere. The transient bioheat applicators can be divided into the following categories: antransfer equation was employed to estimate the temperature nular or concentπc-ππg arrays,²-³ stacxed linear- phased elevation due to the ultrasound power deoosiuon. Then the arrays.⁴ secior-vonex arrays,' tapered linear-phased arrays ^s necrosed tissue volume was predicted bv the isothermal dose cylindncal-section arrav s.⁶ and square-element spherical - volume Computer programs were aiso used to do a parametsection arrays. ^' It was shown that ultrasound phased arravs ric studv 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 hypeπherrrυa cancer therapy Relatively low A. Square-element array with spherical surface frequencies (about 0.5 MHzi and small elemeni size (compared to the wavelength; were employed in these studies The phased array w as made from a spherically curved The studies on square-element sphenc3l-sectιon arrays emtransducer which was divided into small square elements phasized using a large number of elements, up to a lew hunThe wnole arrav was airbac ed It w as constructed from a 298 X. Fan and K. Hynynen: Control ot πtt roxttf tlaaua volume during ultraaouπd curgcry 2M

FiC. I . Con^ήiuration of .r- •αuarc-tn-mt-nι jphfru. wit its cocirdituic ?v»ιcm

spherically curved PZT4 bowl, having a diameter of 100 where μ,_t, is inc absorption coelficieni and Z = oc is the mm. a radius of curvature of 130 mm. 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 siαe Λ 0 3-mm space between the elements was filled with sύicone rubber for electrical and mechanical 0. Inverse technique isolation. Each oi the elements u. as connected to an LC matching circuit to match the impedance to 50 Ω and 0^e The The inverse lecnnioue can be used to calculate ihe ampliarray was driven » un a custom made 16 channel amplifier tude anc ona e sellings irom selecleo control points wnerc (Labiherrmcs. Champaign. Illinois). The phase and amplime αestrec pre ure; \ alues are en en. Defining tude were controlled bv RF signals feeding the amplifiers The driving signals a ere generated by an in-house manufacdS, . tured digital circuit * Tne amplitude and phase of each input signal were dieitaiiv controlled with 3-bιt resolution in arr, men lor M ielα points. expression . l ι can written :r. plitude and 4-bιt resolution in phase Tne numccr of e rr.a'.πx nctaiion a; ents in each rou ano each column were the same so iha' the whole applicator u as a square snape focuseo iransducet Tne configuration of ihe pnased array for the experiments ; • shown in Fie 1

θ. Ultrasound field measurements

The relative pressure amplitude squared distributions i p

were measured in αegassed water using a needle hyαrophone (active spot sue 1 mmj scanned across the focal region. The needle hydrophone as moved bv stepper motors, typically with 0 I -mm steps ... OS the beam. The total acou'iic power (.' = was measured using a radiation torce technique.

Tne eiemems ol matrix H are evaiuatec DV nume .cai inie -

C. Computation of ultrasonic fields e tion usin; eiDirssior (3) Equauon (4 ι has iv .mpoπani fea_tures Firs: if ihe H matrix is calculated ar.c saved for _

Consider an uura<ound phased array with Λ elements αesired ήeid then the pressure field can oe ev aluated using emploveα 10 produce ;m ultrasonic field in a nonanenuaunp Eq ι41 instead of Eq ( 1 1 for a given cxciuuon so rce matrix medium Assume tr.c coordinate system is defined as in Fig V ComDuiation nme can De saved ov using .nis method 1. According to ine ineory of ultrasonic radiators Developed when v ary ing tne excuølion source over the S3mt calculation by O^' N'eil'" and tne principle ot superposition of acoustic \ olurr.; Second :or a oesireα pressure field D2'.:em. i.e.. lor pressure, the preour; held due to tfits transducer can oe a en en τv.ivr. P :ne e ciiation source I car. εs calculated evaluated bv ihe

leιgrι- Sommerfeld iniegr.il The iπte- cral is

O - H ^' P

^ ∑ 'IS., wnerr H' is tne cseudomvt-rse ma x of H. Usually the total numDer ot control points re much le«s than me total number TΛBLZ J. Selected control poirut used lor iπcersc ςjkul.,...nc f<ιr 16 square- where T is the temperature Sit Hfine; . . HhWocaιfS^'H,,^,'^'.£ϊ. etemeni spherical curved phased -rτa> The cπnirυl points are )ιχ- ied aι p, is the density of the tissue, c, is Ihe specific heal of the ^• :•■ 129 mm plane Note die -nιt> fiir » and v are millimrtcrv ■ issue, k, is the thermal conductivity of the tissue. « is the blood perfusion rat. . c_h is tl._ speci ic heal of the blot.i. 7„ is the arterial blood temperature, and @(.τ..v.r) is the acoustic power deposition rate per unit mats The thermal response was simulaied in a homogeneous medium. A surface lem- perature of 37 ^"C on the cube and an initial temperature of 37 *C inside ihe cube were used as boundary and initial conditions for all of ihe computations A numerical finite difference method was employed to solve Eq. (51. Previous studies have shown :hat there are seven! parameters which affect the temperature elevation ^π Temperature elevation has been shown to be almosi mdependeni of the blood perfusion raie for short ultrasound pulses ^]i For this reason ultrasound pulse durations of 1. 5. ard 10 s were used in lhis study The maximum temperature reached in all simulations *dS kepi ihe same (80 "O bv aαιu<.tιπg the input power

F. Thermal r osβ calculation end necrosed tissue volume estimation

The thermal dose calculation was based on the iβchnique suggested

Sapareto and Dewey ' Using this technique, the accumulated thermal dose was calculated at a reference temperature DV numerical integration under different temperature profiles The thermal dose, i e.. equivalent time, at ihe reference temperature can be evaluated by ∑ Λ ^r-"- ^;-»'Λr

(61

where T „, is ine reference temperature. i_s._J. = .-,„„_ir,_t -»- f_c._ei,-_s is the final time. _\ ; is a srηjll time interv al. T_{l t} is ihe avera^ge temperature duπns time r and R is a parameter giver of excιi2tιoπ sources in this case. The H^' matrix can be b fo>— by utilizing sin le v alue decomposition

e control points were selected in a plane at a desired focal position. All the points w re evenly distributed on a 1 0 2 : otherwise circle. Since tne pressure at a control point ii a complex The necrosto tissue volume is estimated bv the isothermal number, both the magnitude and the phase are required 10 dose volume surrounded for 240 min at the reference temperform the inverse calculations. In selecting control points perature of 43 "C Tms lecnmque has oeen found to be t in this manner, it is natural to make ihe magnitudes the same reasonable model for oredicting tissue necroses induced by a for all the points, and the phases either in phase or out of spneπcally cuπ ed transducer phase so that the phase rotates around ihe ceniral axis. The A iwo layered meαium waier-tissue was assumed in the latter setup gives desiructive interference on the central axis simulations The speed of sound and the density w ere 1500 thus, eliminating potential hot spots on tne axis ³ The sc rrvs and 998 kg/m ^'. resoectiv ely for botn media. The attenulected control pomis used in the inverse calculations ar: ation coefficient of tne tissue w as assumed lo be 10 Np/m' grven in Table 1. MHz. The thermal properties of the tissue are giver, in Tabic II.

E. Thermal modeling

An approximate temperature response 10 the power depoIII. RESULTS sition was predicted by ihe transient bioheat transfer equaA. Comparison ot experimental results with tion. The differential equation * _> simulations i- T ά'T - It was necessarv to veπf> the numerical model using tne

^' "M — — - ιι c ,,( J^"- T, ⁾ Q[ .\ . \ .z 16 square-element phased array wnh a spherical surface before relying on the simulations The experimental and simu- ami A. ran ano ιv. nynyn. .oniroi ot nccroseα usau* voium* αurtnø iiraaouπα »u- -y

II 185 C 19- 1' 284 I" 3080- t 5 10 10 15 lated profiles of the pressure amplitude squared at the acous_¬ 209 ^ 21 EC 275 C 2997* I 8 15 I 8 tic focus are displayea in Fig 2. All the data were 3119 2123^J 30-13" I 0 i o I 0 normalized to I bv dividing by the maximum pressure am C-s 3J r ιι« ' 49' 10 i c piirude squarεα tor each curve Figure 2(a) snows the results ι:ι :.: ^> :OJ ' 29.9' I 0 i o

J' 9' I 24 for the case wnh umrorm excitation sources There was good 10 agreemeni oeiween the simulations and experimental results 1515 C 142 J' for the main oeam When the array was excited bv various Case n^Q :° 1 : : .0 phase settings, the simulations reasonably predicted the siαe- 232 " 2232" 3132 I 0 1 c

2 80° ::? _' 222 J'

I 0 i : lobe distributions found in tne experiments [Figs 2(b)— ( 1) The maximum power output measured was at least 1 W for each transducer element

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

The pnased arrav ultrasound transducer used in the ct the pressure a' th: control points had 900° pnase shifts from penments was simulated first. A calculation volume of 3" one point to another (Figs.3tcι and 3(d)] Comparing Figs X30X9O mm' was used for this phased array The axial dis 3icl and 31 d ι wnh F ips 3teι and 3(f>. the Dei* to Deax ois tance from tne center of the transducer to the water-tissue tance was increa'ed wnen the control points on a 125-mm interface uas 70 mm Tne inverse technique results ar; radius circle increased to a 32j-mm radius circ.e However snown fust Tne amplitude and phase settings calculated o_. if tne cor.tioi points were on a 522) mm racius circle, the the inverse technique corresponding to selected control local spot;, were at (he same location as Fiεs 3,cι and 3'dι points arc in Table III irjnh the first lour are shown Lie neid distributions mot Sho n fierei were verv similar ιc Tne contour clots ot Dower deposition calculated for ιhe<e es . i aid j o¹ For six control ooir.ts. s x local SDCI array settings arc αispiayeα in Fig 3 For tour control point' were ootaincc at control locations wι:n cirrerc-.t Dartcm- only one focus was produced which was centered on tn. [Figs 3'.eι and 3ι ι, -.no t o additional focal SDOIS were 2t tne center W hen ihe nu oer ot control ooints increased to

F:c 2 Relancc crenure imc-iituαe vou-ieα dι<ιrιbυιιon _c D luncttor c 4(dι] wa J rr.rr. Thz length was more man douole and tne ιcr,ιl dist nce at ine icou iic locus — I 29 mm> The crσnie *ιt ih- widin was sl.gnπv enlarged compared to the uniform excitadoued line »ιs measured -no mt pronie *ιι t e *ohd line »i^ mul-teύ The aτπpιι*uoe« were H me same r.h p accs lai υπiforrr - '. -n id tion case Tnt computed necosT¹ tissue volume was aoou: 339 mm^' Ftπures 4ιeι and 4(f) snow case II of Table IV for

Fic. 3. Con_tour plou o( po»er dcpojioon lo_t calculated amplitude and pnase semncs. TV -intrr* points useα IO calculate arnplnudcs and phases irom TabI: I were (»). lb) caie 1; icl. (d) case 11; lei. it) case III. (g). th> case v. ans ι u .JJ case Vi iTie ήeures σn ine ιe(ι side a/e a»ιaJ plane distributions^, and on in^e right side are focai plane distributions i - = 1 29 mm) 302 X. Fan and K. Hynyn..,. Control ot necrosed t-aaua voluma during uitraaoαnd i. /rr 302

TΛILE. IV The simplified phase and amplitude sellings for ihe I- sqvart- To understand the effect of radius of curvature on the element sphencillv curved phased array med in the necrosed tissue vαlβrπe necrosed tissue volume, the isothermal doses for ;ue IV of

10-s sonication - united volume was created w tm the total

IV. DISCUSSION vo me of S 10 mm' For the amplitude and phase setting ot case IV in Table IV. tne power deposition tnot snown here The etper.meπtal and simulation results sne ed that ? indicated that mere w ere four strong focai points tdeoth 5-1 16-elcm:n- onased

can offer significant control e- mm i surrounded DV tour small focal points The shortest di - ihe size and snaoc ol ;ne necrosed tissue \ oiume during u' tancc between tne sirong pe-"" ac aoout 2\ Simultaneous trasouno <-urge-y Tne experiments showed that si.cn an aπ^ focusing at these points was performed to enlarge tne heated can oc .cns..--.c;.α arύ n can

enough power for SJΓ volume compared to tne unilorm excitation case The calcj c;ca!

simulated amd experimental ulirasound lateα isothermal dose [Figs Jι ι) and Ah⁾] trom the .eπvocid held outnoj-ions .ereeα reasonably well Tre difterences ture distπoution n.α a maximum lengtn of ?S 5 ran and - mav oe cue :o

en element size and pow er output in tn: maximum widtn ^r 7.2 mm and tne necrosed tissue volume phased arrjc _as v_Zi, as measurement errors Since tne Dial^, M-as ciose to I 4 env to peat. dιs;ance .v a onlv a tew millimeters tne ultrasound

To illustrate me effect of

on necrosea tissue detcctc: T_JI contπoute some measurement errors Tne simu vo me production, tne isotπermal doses for case IN^' of Table lation ooc. accuraieh Dreoicted the locations of sideiooe^c- IV wnh an ultrasound puise αuration l ϋ s are snow n in Figϊ and mam otams except at it over predicted tne magnitude 5taι and 5ι b) The .aiculateo volume w as 60 x60 -^ 90 mm' of the sidclobs * vv.nen aαiacent elements of a onased iπzs for 0 5 MHz. 40x-J0x90 mm tor 1 0 MHz. and 30 x 30x90 are tn c iose to each otner. the mutual coupune between the mm'^' for 1 4 and 2 0 MHz The distance Irom me transducer

c an ertect on the ultrasound f.eld The to tne interface » as kept the same ι ^"O mm i n.ie the f-r d.tterence ct' -ctn tie sidelooes in tne experiments and the ouencv was varied Tne ratio ot the necrosed tissue length :^•

be due to the lack of such crosstalk in ine wiotn Aas almost independent of r eαuenc> Tie size in calculation; creased as the treque,.-v decreased Tne shapes were simuat The _«uo>. e results nroved that tne large element spn^π- for treαuencies cf I lc 2 MHz calh cu^ e phaseα array can enlarge tne necrosed tissue 0- Λ-fβnano R Hynynen: Coπτrol oι πc eβo tissue volume ounncj uitrasouno surgery 4UJ.

Z-Distance (cm)

Fic.4 Isochcrm-J doses (or me ur.iicnn e«c,....on case lal .bl and ihe .mplιt-dr and phase seu c n Table ^ ^■<:^{■ i}d, case I ,e> '" «'- » 'f' ^'^ IU and (0. I_jl case IV The n.ure: en .he lefl side arc u.al plane disinouuons and on the n_fh. s.αe are local Dim αitinbui.on, b> := I .» **""•«> - ' ;- mm. ω .« 1-5 mm ihl := » mm. and ι_j, ;= 124 mm] The uurasouno pulse duranons used ,n me inermal oosc calculaucns sc«e lϋ s _lSol,dl.neι . > (dish-U- line), and J s (doncd line y* X. Fan and K. Hynyn^βi omrol ol necros.O tistu* volume during ultroscH-iut- u' >

S 10 U 12 13 li S Iβ -2 - 1 5 -1 -0-5 o oa i u i

Z-Distance (cm) - Distance (cm)

FiO 5 The t^jothermal doic lor cue IV ot^' Tabl (c ) Ed¹ The uir ounα DUIW ύuritioπ used in e thermal dose calculation* was 10 s }2 —.m icr raon ot eun i ure of 20C 160. 130 and SO mm resDecnvelv (_a), l t is

volume. The phased array can also enlarge the necrosed tisThe l o eiemen: pr.ased arrav can generate tcur foca sue volume in only one direction at a time, if desired. It is points w nn 3 near, to o-_ t distance as short as tw o

important to be aole to control the focal spot size so thai i lengths T ne maximum Distance between the losest peans is large rumor could be treated in a reasonable time. The con limned IO arci tour w avelengtns When the leur contro, smjctioπ of the whole system was relatively simple due to points were oi a circle ot radiui 5 25 mm, tne pnase mer the small nu oer of elements The 16-elemc phased »rτa> mem between .Λiacent elements exceeded TΓ for mis arrav can aiso shift the focus off the central axis or move the focus Therrtore n .s Dhvsicallv imrcssible to use tne phase ano along the central axis. However the distance is limited to amplitude seii.ne obtained b\ inv erse techniques to proαuce - 1 5 mm laterally and 14 mm in the axial direction In movthe four tocaj spots at :ne control points Altnough Eq 1 1 ing the focus alone the central axis, the array is similar to _ represents an unoerdete minα svstem. it producea a solution concentnc-πne array witn two nngs. Theoretically, tne maxi at the control DOIM*. V> nsπ multiπle loo art separated bv mure phase increment between adiacent elements if - distances large' tr.an A 6 mm. the necrosed tissue volume Therefore, the maximum possible pnase difference between becomes a teu. individual smai er

hich arc no' the smallest ana largest phases is IT w hen moving the focuc unned lor snoπ ultrasound ouise durations along tne central axis From geometric considerations, this Bv decreasing the freαuencv. the necrosed tissue volune phase difference Droouces displacements along the central can be enlarged oecause the focal spot increases oue to the axis of up to 23 mm f lO mm closer, 13 mm oeeper) In increased w avelength But the ratio of tne necrosed tissue shifting the focus sideways, it is similar to a cyiindπcal- length to the u idtn w as xept almost the same As tne raoius secnon array with four elements. The maximum possible ot curvature is increased the necrosed tissue volume can also pnase difference between the smallest and largest pnases is be enlarred The necrosed tissue voiume is increased maιnl> 3 w for shifting. the focus sideways Geometrically, mis phase in the aua. o:rectιon Tnis aereeS u itn previous exoenence difference shifts the tocus 3 A mm laterally How ever, the using single element scneπcallv curved transoucers simulations snowed thai the phase increment between the Phased

reαu-.re more input pow er man similar adjacent elements shot¹' ' be wss ιh»n -12 to generate a -'nele tocuseα traπsrtucrrs to reac . tne same temperature ainsle strongly tocused ultrasound field level for tne same ultrasonic pulse duration due to increased

Z-Distance (cm)

ACKNOWLEDGMENT

Fio 6. Contour plots of po» er deposition for v inous amplitude ani pnase settings tai i; ihe uniform excitation case Tie cv.a) distance Irom ihe This sludy w as supported by NCI Grant No CA 4662^* transducer to ihe focus »as ι a > 1.9 — m Ibi I-- mrr. " f I jb mm. and ιoι 129 mm. shifted 1.5 mm

¹ L nιc er<ιtv of *rι

'K

The (f reehold lor therrπalK cicπhcani cavitatton in dec s lhiRh muscle in ι ι . J L nrasounύ Med Bid 1 " 1 57- 169 ( 199 ) focal spot size. The test array measurements showed thai "J P Do-Huu an i P Hanemzr.r, Annular arrav transducer tor Cee, therequired power levels can be generated in practice at least acoustic nvpenhermij t_ Uπ»orιcs Si m Proc IEEE-S I CH 1 1*9-

^">^» and 10-s exposures and thus, the proposed technique is 705-710 (19B I 1 iι_-sιble 'K B. Och«lιr« P J Benkesei L A Fn2-ell. and C A Cain ^c\π ultrasonic phased arrav applicaior (or hv peπnermia. IEEE Trans Sonics

In Order to obtain the desired field pattern, several field Uluason . Si'-? I .*.6-531 ι l9β-; ι control points axe necessary to perform the inverse calcula*C Cain and S Umemura Conceninc-nnr; and se tor-_*eπeι phased tions. However, the number of utiliiable field control points arrav applicators lor ultrasound nvpcrrtherrnia. IEEE Trans Microwave

Theon Tech MTT-34 x;-; i < 19S6> is limited by the number of phased array elements. Hence the 'P. J. Benkeser L \ Fπi-ell K B O hrltrce and C A Cain. A tapered number of field patterns that produce significantly different phased arrav ultrasound irancdυ cr for hv perthermta treatment. IEEE shapes or sizes of the measured volume is limited. Although Trans l ltrason Ferroelec Frcq Conir L FC 4 4J6--tf 3 09S7 ' only few amplitude and phase settings were presented in this 'E S Ebbini. S Umemura i lboim and C A Cam. A ev ndncal- section ultrasound phiseo-arrav applicator for hvpeπhemia cancer study, the results provided information which could aid in t erapv" IEEE Trans Ultrason Ferroelectr Fret} Conir 35 561 -572 planning ultrasonic surgery using a phased array system n • 19fcS) a small number of elements. These results have not been E Ebbini and C A Ca n A sphencal seαion ulirasound phased arric optimized, nevertheless they illustrate the feasibility and applicator for deep localized n-.peπherm.a. ^■ IEEE Trans -t'omed. tnf

BME-.'S 634 -6-3 D99 I I range of focal sue and shape obtainable with a 16-element 'E. Ebbini and C A Cam. Multiple-locus ultrasound pha<ed--rrιv pn array. The calculations and experiments can be used to obtain tern synthesis Opnmal dm mj-cipnul distnbutions for hvpeπhermi* a set of different focal spots for a giv en arrav This limited IEEE Trins Ultrason Ferrociectr F(-q Contr 36. 540-5JS 11989, *M. Buchanan and Hvπvneπ Design and eipeπmental evaluation of number of focal spots can then be utilized to optimally cover iniracavitarv ultrasound phased array s stems lor hvpenhermia." IEEE the target volume in minimum time. The minimization of the Traπc Biomed Enc BME-- I I I 7S- 1 187 U99*l duration of the treatment is an important factor for control^J0H T O Neil Tneorv oi locusinp radiators ^" J Acoust Soc Am 21 ling the cost of the procedure, especially wnen MRI is used 516-5:6 ( 19491 "X Fan and K Hvnv πen The effect of vi -ve reflection and refraction aι to guide and monitor tne treatment '" If more control over soli tissue interfaces duπnc ultrasound hvpeπhennia treatments.- J the field pattern is required then an array uh a larger numAcousi Soc Am 91 1727- 1736 ( 1992! ber of elements is required. Such an array could also be used '"X. Fan and K Hvnynen Ths {fleets ui curved ussue lasers on the 306 X Fan and K Hynynen: Control of necrosed *lseue volume during ultrasound surgery 308 power deposition pattern, of therapeutic ultraiounu beaπu," Med Phys rameters on high lemperature ultrasound hypeπhermia.'- Ultraiound Med

21. 25-34 (1994). Biol It. 409-420 (1990) ¹ 'C. Damunou and K. Hynynen. " The effeci of vinous physical param- ^l5S A Sapveio and W C. Dewey, rheπnal dose deιerrnιn__ιon in cuicer eiers on the size and shape of necrosed tissue volume duππg ultrasound therapy. Ini. J. Radial Oncol Biol. Phys -10, 787-800 (1984) surgery." J Acoust Soc Am. 9S. 1641 - 1649 ( 19941. "K Hynynen. A . DaΛa_anlι. £ ϋnger and 1. Schenck. "MRl-puided B E Billard. H πynen. and R B Roem.r. Effects of ph) sical panomπvasn e ultrasound surgery ' Med Phvs 20 107- 1 1 j (1993)

Claims

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 with respect to one another.

2. An apparatus according to claim 1, wherein

A. the 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,

B. so that the ultrasound generated by the selected transducers arrive substantially in phase with one another at the selected region.

3. 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 90° of one another at the selected region.

4. 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 45° of one another at the selected 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.