WO1998007373A9 - Methodes et dispositifs pour administrer un traitement par ultrasons non invasif au cerveau a travers une boite cranienne intacte - Google Patents

Methodes et dispositifs pour administrer un traitement par ultrasons non invasif au cerveau a travers une boite cranienne intacte

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Publication number
WO1998007373A9
WO1998007373A9 PCT/US1997/014760 US9714760W WO9807373A9 WO 1998007373 A9 WO1998007373 A9 WO 1998007373A9 US 9714760 W US9714760 W US 9714760W WO 9807373 A9 WO9807373 A9 WO 9807373A9
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WIPO (PCT)
Prior art keywords
ultrasound
transducers
skull
selected region
mhz
Prior art date
Application number
PCT/US1997/014760
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English (en)
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WO1998007373A1 (fr
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Publication date
Priority claimed from US08/711,289 external-priority patent/US5752515A/en
Application filed filed Critical
Priority to AU42333/97A priority Critical patent/AU4233397A/en
Publication of WO1998007373A1 publication Critical patent/WO1998007373A1/fr
Publication of WO1998007373A9 publication Critical patent/WO1998007373A9/fr

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  • 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.
  • 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.
  • 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.
  • the invention provides in one aspect methods and apparatus for delivery of cavitating ultrasound to the brain, without requiring removal of portions of the skull.
  • 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.
  • 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.
  • 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.
  • FIGS 5 A and 5B illustrate the effect of applying ultrasound in accordance with the invention to brain tissue.
  • Figures 6 A and 6B illustrate phase errors measured at the focus of ultrasound transducer arrays with a piece of skull in front of the transducers.
  • Figures 7A-7C illustrate the pressure amplitude profiles across the focus of an ultrasound transducer phased array in water, through the bone, and through the bone with phase correction.
  • Figure 8 illustrates the pressure amplitude distribution along the central axis of an ultrasound transducer array without and with the phase correction.
  • Figures 9A-9C illustrate the ultrasound pressure amplitude distribution measured across the focus of an ultrasound phased array in water, through skull without phase correction, and through skull with phase correction.
  • Figure 10 depicts an embodiment of the invention for delivery of cavitating ultrasound to a patient's brain through the skull using a multi-element transducer array.
  • Figure 11 depicts a method for delivery of cavitating ultrasound to a patient's brain through the skull using a transducer array.
  • 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.
  • 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 cm 2 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
  • 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.
  • RF radio frequency
  • 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).
  • 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.
  • 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 ⁇ ,, 2 , ⁇ 3 , etc., as shown in the drawing.
  • These phase shift factors ⁇ ,, ⁇ 2 , 3 , etc. can be pre-stored in the channels of driving system 18 or, preferably, generated by a controller 20.
  • That controller 20 can be a general purpose, or special purpose, digital data processor programmed in a conventional manner in order to generate and apply phase shift factors in accord with the teachings hereof.
  • phase shift factors serve two purposes.
  • the first is to steer the composite ultrasound beam generated by transducer array 12 so that it is focused on a desired region within the patient's brain.
  • the component of each phase shift factor associated with steering is computed in the manner known in the art for steering phased arrays. See, for example, Buchanan et al, "Intracavitary Ultrasound Phased Array System," IEEE Transactions Biomedical Engineering, v. 41, pp. 1178- 1187, the teachings of which are incorporated herein by reference.
  • Array steering, or focusing is particularly discussed in that article, for example, at pages 1179-1181 and, more particularly, in the section entitled “Focusing Techniques,” the teachings of which are incorporated herein by reference.
  • each phase shift factor ⁇ ,, ⁇ 2 , ⁇ 3 , etc. compensates for phase distortion effected by the skull in the ultrasound ouput by each transducer.
  • 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.
  • 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.
  • phase corrections that constitute the aforementioned second component of each phase shift factor can be determined a number of ways.
  • that component is determined from measurements of the thickness of the patient's skull under each transducer in array 12.
  • Such skull thickness measurements can be made using conventional imaging techniques, such as computed tomography (CT) or magnetic resonance imaging (MRI).
  • 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.
  • 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.
  • 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.
  • a driving system would, of course, require modification in accord with the teachings hereof in order to incorporate phase shift factors ⁇ ,, ⁇ 2 , ⁇ 3 , etc., having first and second components as described herein and above.
  • 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.
  • 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.
  • 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.
  • 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.
  • the array is excited, e.g., by control and driving elements 18-22, to focus ultrasound in the patient's head.
  • 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 mm 3 - 1 cm 3 , or larger.
  • 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.
  • step 26 is not utilized.
  • 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.
  • 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:
  • ⁇ q> ⁇ P 2 / pv
  • the ultrasound beam 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 (P 2 ) 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:
  • 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:
  • ⁇ is the insertion loss of skull
  • ⁇ is the amplitude attenuation coefficient
  • f is the frequency
  • 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.
  • 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.
  • 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.
  • 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 Wcnr 2 and 2000 Wcnr 2 at 1 ms and 1 s exposures, respectively.
  • 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 are 1800° C and 3.6° C.
  • 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.
  • ultrasound was generated using a one-element transducer array having a spherically curved 10 cm diameter piezoelectric ceramic (PZT4) bowl mounted in a plastic holder using silicon rubber.
  • the ceramic had silver or gold electrodes both on the front and back surface.
  • the electrodes were attached to a coaxial cable that was connected to a LC matching network that matched the electrical impedance of the transducer and the cable to the RF amplifier output impedance of 50 ohm and zero phase.
  • the matching circuit was connected to an RF- amplifier (both ENI A240L and A500 were used in the tests).
  • the RF signal was generated by a signal generator (Stanford Research Systems, Model DS345).
  • the ultrasound pressure wave distributions were measured using needle hydrophones (spot diameter 0.5 and 1 mm) and an amplifier (Precision Acoustics Ltd).
  • the amplified signal was measured and stored by a oscilloscope (Tektronix, model 243 IL ).
  • the hydrophone was moved by stepper motors in three dimensions under computer control.
  • the pressure amplitudes measured by the oscilloscope were stored by the computer for each location.
  • a piece of human skull (top part of the head: front to back 18 cm and maximum width 12 cm) was obtained and fixed in formaldehyde.
  • the acoustic properties of formaldehyde fixed skull and a fresh skull are almost identical.
  • the ultrasound transducer under test was positioned in a water tank the walls and bottom of which were covered by rubber mats to reduce ultrasound reflections.
  • the tank was filled with degassed deionized water.
  • the hydrophone was connected to the scanning frame, and positioned at the focus of the ultrasound field.
  • 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.
  • 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.
  • 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.
  • the Figure is a T2-weighted fast spin echo image across the brain.
  • the arrow in the Figure shows tissue damage at the focal point of the transducer.
  • the skull window on the top of the head is facing down and, thus, the ultrasound beam propagated from bottom up.
  • Figure 4B is identical to Figure 4A, except insofar as it shows the results where the above sonication was repeated four times.
  • FIG. 3 For embodiments of the invention, these embodiments utilize multi-element phase arrays of the types illustrated in Figures 3 and 10, in lieu of a single transducer.
  • phase of the ultrasound wave 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 cm 3 .
  • 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 cm 2 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.
  • 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.
  • FIG. 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.
  • phase distortion caused by the skull 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.
  • FIG. 7 A 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.
  • 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 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.
  • 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.
  • results 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/cm 2 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.
  • 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.
  • each wave is adding to the pressure wave at the focus.
  • 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.
  • 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.
  • full utilization of the skull surface may provide marginally adequate geometric gains to overcome the skull heating problem.
  • the focal brain tissue thermal therapy seems feasible although not as likely as utilization of cavitation effects.
  • 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.
  • 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.
  • 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.
  • 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.
  • INTRACAVITARY ultrasound arrays offer an attractive proposed or built for hyp ⁇ henmic purposes. These include means of inducing local hyperthermia in deep-seated tuUmemura and Cain ' s sector-vortex and concentric ring applimors located near body cavities.
  • the radiators By locating the radiators as cators
  • Each of these arrays is composed of the acoustic window by bone or gas, or simply the inability anywhere from 16 to 64 individual elements and operates at to attain adequate energy penetration, can be avoided.
  • Early frequencies between 500 kHz and 750 kHz. While these arrays results using multielement, nonfocuscd arrays of half-cylinder show significant potential, they ate meant to be used in external transducers operating at 1.6 MHz suggest that such arrays can applications and therefore are unsuitable for intracavitary use be clinically useful in the treatment of prostate cancer ( 1 ). in their reported configurations.
  • the proximity of the prostate to the rectum wall makes it Previously, a theoretical study on the feasibility of intracava good candidate for heating using intracavitary ultrasound itary phased arrays using half-cylinder elements had been done radiators. Since the prostate is located very near the anus and by Diederich and Hynynen
  • the prostate is one of would be composed of 30 half-cylinder elements with 2.5 the most easily accessible tumor sites, and one that affects a mm center-to-center element spacing operating at 500 kHz.
  • the amplifiers are based on a switching MOSFET design and are designed around International Rectifier's IR2110 dual
  • a Array Cons ⁇ itction MOSFET driver Each of the amplifiers is capable of deliv ⁇
  • a 64-element array was constructed using half-cylinder ering up to 16 of RF power at 500 kHz into a 50-W load transducers operating in their resonant radial mode at 500 kHz from each channel or a total output power of about 850 W.
  • the array was made by slicing washer-shaped elements with The amplifiers convert digital logic input signals into high a diamond wire saw (Laser Technologies. North Hollywood, power sine waves while preserving the phase of the input CA) from 15 mm O D by 30-mro long cylinders of P2T- Nignal
  • the amplitude of the output signal is controlled by 4 material (EDO.
  • the cycle is the percent of "on" time of the input signal per clock stack of elements was then cut in half along the axis of the cycle) ts a corresponding decrease in the amplitude of the cylinder and the two half-cylinder sections glued together to output signal. form the full array
  • the array was bonded to a brass shell Since the amplifiers require digital input signals, the phase to form the complete applicator, as shown in Fig. 1 Wires shifting and duty-cycle control is implemented using digital were soldered to the inside wall of each array clement that ,ou ⁇ ters ( 10], [ 1 1]. These circuits provide 22.5° phase shift extended the length of the shell to the handle where they were resolution from 0-360°.
  • phased arrays are the by power meters that measure both the forward and reflected increased complexity of the driving equipment. Due to a lack RF power [13]
  • the power meters which were also designed IEEE TRANSACTIONS ON BIO EDfCAU'ENifc. .eERINOrVOL -tf. NO ,” 12.TJE t BER 1994
  • n is the particle velocity normal to thr surface of the source
  • r m - r n is the distance between a and n z is the number of sources in the 2-d ⁇ rect ⁇ on
  • n $ is the control point (r m ) and the surface of a source (r n )
  • m i, number of sources in the theta direction
  • P s is the pressure
  • P sa is the pressure amplitude at the
  • the number of elements intracavitary hyperthermia uses, the number of elements in
  • the single focus case is the simplest form of focusing that points (M) This leads to an underdetermined system of can be done with a phased array
  • the single focus is produced equations with an infinite number of solutions
  • the minimum by setting the phases of the driving signals so that constructive norm solution ( ⁇ ) can be determined by using a least squares interference occurs at the desired focal position
  • the phase of approximation of (5) each of the driving signals for an array with N elements can be calculated from the differences in the path lengths ⁇ f, between u - H' ⁇ HH'*) 'j (6) each array element and the focal position by where II" is the complex conjugate of H
  • all of the array elements contribute to all of the foci, unlike
  • the other technique multiple focusing, simultaneously produces more than one focus within the target volume
  • the drivB Ultrasound Field Measurements ing signals necessary to produce multiple foci were calculated
  • the ultrasound fields were mapped in a tank of degassed, using two techniques split focusing and the pseudo-inverse deionized water by mechanically scanning a thermocouple
  • the ther ⁇ array was divided into subarrays, each of which produce a mocouple was positioned by a three-axis computer controlled single focus m the same manner as previously descnbed
  • the applicator was mounted on a rotational pseudo-inverse method, developed by Ebbini and Cam [15], device that allowed measurements to be made in a radial arc uses a series of control points that represent the magnitude around the array by rotating the array Measurements were of the ultrasound field at given points
  • Alcohol fixed canine kidneys were used as phantoms for studying the heating characteristics of the array.
  • the kidneys had previously been prepared as described by Holmes et al. [18], and were rehydrated prior to use.
  • the experiments were conducted at room temperature using degassed, deionized water as the perfusate.
  • a metering pump (Fluid-Metering Inc., RHICKC, Oyster Bay, NY) connected to the renal artery circulated water through the kidney while the renal vein was allowed to drain into the tank.
  • the kidney was held in place by gently sandwiching it between two PVC membranes mounted Pu ⁇ to a Plexiglas frame.
  • the applicator was firmly clamped to the I ⁇ fatar Bach frame to maintain a fixed distance between the surfaces of the Fig. 3.
  • Fig. 3 shows a diagram of the experimental setup.
  • Temperatures were measured in the kidney using either 20 - seven sensor probes sutured in place perpendicular to the array through the focal region, or by one or two single sensor thermocouples pulled along a path parallel to the array.
  • the kidney was exposed to ultrasound for a total of 10 min with temperature measurements occurring every 30 s.
  • power to the array was disrupted for approximately four seconds (one second prior to the first reading, and up to three seconds to read all of the thermocouples). Since this technique does not give very fine spatial detail of the temperature distributions, the pull-back technique was more frequently used.
  • the pull-back experiments were conducted by pulling one or two single uncoated thermocouples (0.05-mm w ire) by a 10 15 20 computer controlled stepper motor along a track parallel to Electrical Power (W) the array through the kidney tissue.
  • the temperatures were Fig. 4.
  • Acoustical output power as a function of electrical input power for measured every 1 mm and the average of three readings was a single 30-mm. full cylinder element, and arrays made up of 16 2.2-mm and recorded.
  • a baseline tem1.5- n ⁇ elements with a 0.3-mm insulator between elcmenis. perature was established along the path of the thermocouples.
  • the kidney was exposed to ultrasound for 20 min to allow (width ⁇ wavelength), which would produce nearly spherical the temperatures to reach steady-state before the temperature wavefronts. The collimation is necessary to assure the waves profiles were measured, The difference between the baseline reflected by the reflector are normally incident to the surface and final measurement was used to calculate the temperature of the absorber. rise. The kidney was allowed to cool to room temperature The results of these experiments are shown in Fig. 4, None (typically, 30 min) before to the next experiment was started. of the arrays exhibited very high efficiencies despite the low
  • thermocouples were mainly located in the medulla of operating frequency. Though not shown here, some results the kidney. The steady-state temperatures in the medulla have were verified using calorimetric techniques. While the 30-mm been shown to be a strong function of the flow into the kidney long full cylinder element had an efficiency of 71 %, the arrays [19]. The flow values were kept relatively low in order to exhibited efficiencies of 27% and 17% for the 2.2-mm and 1.5- simulate the perfusion in the prostate [20], [21 ]. m arrays, respectively. These efficiencies indicate that about 80% of the electrical power delivered to the 1.5-mm wide
  • FIG. 5 Field plots in water of two lo-elemem arrays with l ft-mm (a) and 2 5-mm (b) element center-to-cemor spacing and their corresponding simulation results (to and (d)> Both arrays were focused 30 mm tro ⁇ i the .i ⁇ ay surt-cc and 15 mm from the central axis
  • the 2.5-mm center-to-center element spacing (Fig. 5(b) and (d)) acoustic field was measured as a function of rotation angle Since such a large grating lobe not only reduces the power along a line by fixing the position of the thermocouple in of the focus but also could cause heating in unwanted areas, the focal region and rotating the array about its axis.
  • the arrays must be designed to minimize grating lobe formation. acoustical intensity is not at all uniform and, in fact, vanes
  • the array with 1.8-mm center-to-center spacing produced as much as 50% before tapering off at the edges.
  • this array is capable of focusing to on either side of the central axis of the array (20 mm apart) about 35° from the central axis of the array without producing using the 64-element array.
  • the pseudo-inverse significant grating lobes. technique produces sharper foci than those produced by the IEEE TRANSACTIONS ON BIOMEPIC ⁇ I, frfG'lr- "YlP 4l spanIN ⁇ l 11. DBCEMBB ⁇ wl
  • thermocouples were located 8 mm and 15 mm from the surface of the kidney nearest to the array
  • Fig 7 Single dimensional radial held plot made within the focus as a function of the ⁇ ngle of rotation of the array
  • the center of the arrav arc (35 mm and 42 mm from the surface of the array).
  • The is defined as the zero point scanning technique produced a narrower profile than the two multiple focusing techniques and the distribution lacks the temperature drop between foci
  • the two multiple focusing split focusing method This occurs because the pseudo-i crse techniques produce virtually identical profiles, although the technique utilizes the entire length of the array allowing it to pseudo-inverse technique does not produce as large of a produce sharper foci than the spin focusing technique which, temperature ⁇ se on the deeper thermocouple (beyond the since it divides the array into two subarray , effectively uses locus) as does the spin focusing technique.
  • Fig. 9(a) shows the temperature ⁇ se ver us time along a fixed seven-sensor thermocouple probe located perpendicular V DISCUSSION AND SUMMARY to the array
  • the distances marked denote the distance of An intracavitary ultrasound phased array composed of half- the thermocouple from the edge of the kidney nearest to cylinder transducer elements has been constructed for inducing the array.
  • the array, with 32 active elements, was focused hyperthermia in the prostate via the rectum.
  • a 64-channel along its central axis and 10 mm into the kidney has also been designed and constructed to two active elements were used since the array length in this drive the phased array.
  • the array is capable of producing focused fields as well as was perfused at a rate of 2.9 kg m -J s ⁇ l
  • the arrav is currently is only 35 mm thick and thus, the temperatures close to the capable of producing a 12°C temperature ⁇ se in a perfused surfaces of the kidney are dominated by cooling caused by phantom using a stationary focus, and smaller temperature BUCHANAN AND HYNYNEN INTRACAVITARY ULTRASOUND PHASED ARRAY SYSTEM
  • Fig. 8 Acoustical field plots produced u» ⁇ ne the pseudo-mierse (a) and split focusinj (b) techniques and the related simulated results ((c) and (d)).
  • the foci were produced 40 mm deep, and 10 mm on either side of the central axis of the arr-y
  • phased arrays show considerable temperature were caused by a variety of problems, including potential for improvement over currently used intracavitary morphological differences in kidneys and the location of the ultrasound hyperthermia system. thermocouples within the kidneys, the inefficiency of the array, While phased arrays allow significantly more control over and the open loop manner in which the array was ooerated the acoustical field, the current design using half-cylinder
  • the electrical efficiency of the array would usually drop radiators still lacks control in the angular direction (around the considerably during the first experiment, and dunng each subarc of the array). Additionally, since the cylindrical radiators sequent experiment due to changes in the electrical impedance do not have uniform angular intensities, the angular heating of the array elements. Two primary factors were responsible pattern is somewhat degraded, though thermal conducuon will for the observed changes in the electrical impedance of the probably smooth the resulting temperature distribution. Similar array elements: The first was a thermally induced impedance fluctuations in the angular field distributions have been shown drift caused by the array self-heating during sonication.
  • thermocouple locations indicate ihe (b) from the surface of ihe kidney with the array positioned 27 mm from the distance from the surface of Ihe kidney (a)
  • the axial distribution measured kidney surface 12 mm from the surface of the kidney using a pull-back thermocouple (b) in a separate expc ⁇ ment minimized grating lobe formation, higher frequencies would pie-shaped subelements and driving each subelement indepenincrease the power absorption in tissues Since most tumor dently.
  • Focused high-power ultrasound beams are well suited for noninvasive local destruction of deep target volumes
  • ,h ⁇ gh frequencies (1-5 MHz) are used ultrasonic surgery
  • the focal spots generated by sharply focused transducers become so small that only small tumors can be treated in a reasonable time.
  • Phased array ultrasound transducers can be employed to electronically scan a focal spot or ⁇ o produce multiple foci in the desired region to increase the treated volume.
  • phased array could control the necrosed tissue volume by using closely spaced multiple foci
  • the phased array can also be used to enlarge a necrosed tissue volume in only one direction at a time, i e . lateral or longitudinal.
  • the spherically curved 16 square-element phased array can produce useful results by varying the phase and amplitude setting Four focal points can be easily generated with a distance of two or four wavelengths between the two closest peaks
  • the maximum necrosed tissue volume generated by the array can be up to sixteen times the volume induced by a similar spherical transducer Therefore the treatment time could be reduced compared wuh single transducer treatment.
  • Phased array applicators were introduced to ultrasound given array. Several simplified amplitude and phase settings hyperthermia cancer therapy in the early 1980's. Dunng the based on the calculated amplitudes and phases were empast decade, many efforts have been made to investigate the ployed for ultrasound field calculations. These d ⁇ ving signal advantages of phased arrays in hyperthermia. and several sets can be utilized, when different focal spot sizes are rephased array applicators have been developed. Phased array quired for the array proposed hexe. The transient bioheat applicators can be divided into the following categones: antransfer equation was employed to estimate the temperature nular or c ⁇ ncent ⁇ c- ⁇ ng arrays, 2-3 stacked linear-phased elevation due to the ultrasound power deposition.
  • oon ⁇ . ⁇ J ⁇ ' + ' ,,) is the complex excitation source of the nth element with amplitude AZA and phase ⁇ profession , r m- is the distance from : ⁇ point (xicide ,y n ,z r ) on ihe nth element to the field point of interest (x m ,y m ,z m ), S terminate is the area of the nth element.
  • the power deposition for the desired volume is given by
  • the bowl was cut into 16 elements, each with a length of 20 mm per side A 0.3-mm space between the elements was filled with silicone rubber for electrical and mechanical D. Inverse technique isolation.
  • Each of the elements was connected to an LC matching circuit to match the impedance to 50 i and 0°
  • the inverse technique can be used to calculate the ampliarray was driven ⁇ * ⁇ th a custom made 16 channel amplifier tude and phase settings from selected control points where (Labihermics, Champaign. Illinois). The phase and amplithe desired pressure values are given.
  • T MNI were measured in degassed water using a needle hydrophone (active spot size 1 mm) scanned across the focal region.
  • the needle hydrophone u as moved by stepper motors, typically with 0 I -mm steps across the beam
  • the total acoustic power was measured using a radiation force technique.
  • Equation (4) has two important features. First, if the H matnx is calculated and saved for a
  • the excitation source U can be calculated evaluated by the ' Ra> le ⁇ gh-Sommerfeld integral
  • W 4 is the pseudomvcrse matrix of H.
  • T is the temperature iiat
  • Phase blood perfusion rat, . c, is tl... specific heal of the bloc.l.
  • T raw is the arterial blood temperature
  • Q(x.y.;) is the acous ⁇
  • T Environment ( is the reference temperature.
  • condiment ng +f C ooi ⁇ - > _ is the final time, ⁇ ; is a small time interval.
  • T ⁇ f is the aver- as;e temperature during time ⁇ r.
  • R is a parameter given of excitation sources in this case.
  • the speed of sound and the densify were 1500 thus, eliminating potential hot spots on the axis. 3
  • the attenulected control points used in the inverse calculations are ation coefficient of the tissue was assumed to be 10 Np/m given in Table I. MHz.
  • the thermal properties of the tissue are given in Table II.
  • T .BLE II Thermal properties of ihe tissue used in >he simulations.
  • TABLE III Phase and amplitude setungs calculated with the inverse technique for the 16 square-element spherically curved phased array.
  • Figure 2(a) shows the results III 302 3° 204.9° 294 9° 214.3* 1 0 1 0 1.0 1.0 .3° for the case with uniform excitation sources. There was good ⁇ 11 9° 124 3° 32 131.9° 1 0 1 0 1 0 1.0 agreement between the simulations and experimental results 158 0° 142 4° 59.2° 68.0° 1 1 I 2 1.0 1.1 for the main beam.
  • the simulations reasonably predicted the side- IV 232 4- 223 2° 313 2° 329.2° 1 2 1 0 1.0 1 0
  • the isothermal doses for tissue necrosis for five amplitude and phase settings are displayed in Fig. 4.
  • the shape of the necrosed tissue volume was close to an ellipsoid.
  • the calculated volume was about 85 mm 3 , measuring 3 mm laterally and 18 mm longitudinally.
  • Isothermal doses are also shown m the same figures for the ultrasound pulse durations of 5 and 1 s. where the input power was adjusted so that the maximum tempera ⁇
  • X-Distance (mm) ture was kept at 80 °C.
  • the necrosed tissue length [Fig. 4(c)] was 42 mm, and the width [Fig.
  • FIG. 6(d) shows the phased array which the necrosed tissue length and width were 20.8 and focus shifted 1.5 mm off the central axis. When larger dis5.1 rnm, respectively. The length was slightly longer and the placements were attempted, the phase increment between adwidth was almost double compared with the uniform excitajacent elements exceeded ⁇ r/2. Distributions generated by tion case.
  • the power ranged from 30 to 250 W.
  • the volume was about the same size as the volume in the contemplatnput power could be as high as 900 W, depending on the form excitation case for the 1- and 5-s sonications.
  • 10-s sonication a united volume was created with the total volume of 810 mm 3 .
  • the power deposition (not shown here) The experimental and simulation results showed that a indicated that there were four strong focal points (depth 54 16-element phased array can offer significant control over mm) surrounded by four small focal points.
  • the isothermal doses for case IV of Table lation model accurately predicted the locations of sidelobes IV with an ultrasound pulse duration 10 s are shown in Figs. and main beams except that it over predicted the magnitude 5(a) and 5(b).
  • the calculated volume was 60X60X90 mm 3 of the sidelobes.
  • 40X40X90 mm 3 for 1.0 MHz. and 30X3OX9O are very close to each other, the mutual coupling between the mm 3 for 1.4 and 2.0 MHz. The distance from the transducer elements would have an effect on the ultrasound field.
  • the phased array can also enlarge the necrosed tis
  • the 16-element phased array can generate four focal sue volume in only one direction at a time, if desired. It is points with a peak to peak distance as short as two waveimpo ⁇ ant to be able to control the focal spot size so that a lengths. The maximum distance between the closest peaks is large tumor could be treated in a reasonable time. The conlimited to about four wavelengths. When the four control struction of the whole system was relatively simple due to points were on a circle of radius 5.25 mm, the phase increthe small number of elements. The 16-element phased array ment between adjacent elements exceeded ⁇ for this array.
  • the necrosed tissue volume Therefore, the maximum possible phase difference between becomes a few individual smaller volumes, which arc not the smallest and largest phases is it when moving the focus united for short ultrasound pulse durations along the central axis. From geometric considerations, this By decreasing the frequency, the necrosed tissue volume phase difference produces displacements along the central can be enlarged because the focal spot increases due to the axis of up to 23 mm (10 mm closer, 13 mm deeper). In increased wavelength. But the ratio of the necrosed tissue shifting the focus sideways, it is similar to a cylindrical- length to the width was kept almost the same. As the radius section array with four elements. The maximum possible of curvature is increased, the necrosed tissue volume can also phase difference between the smallest and largest phases is be enlarged.
  • the necrosed tissue volume is increased mainly 3 ⁇ r for shifting, ⁇ he focus sideways. Geometrically, this phase in the axial direction. This agrees with previous expenence difference shifts the focus 3.4 mm laterally.
  • the using single element spherically curved transducers. simulations showed that the phase increment between the Phased arrays require more input power than similar adjacent elements shoi'" * be less th»n ⁇ /2 to generate a "mgle focused tra ⁇ sriiicrrs to reacu the same temperature single strongly focused ultrasound Meld level for the same ultrasonic pulse duration due to increased

Abstract

Méthodes et dispositif pour administrer des ultrasons au cerveau sans qu'il soit nécessaire d'enlever des parties de la boite crânienne, faisant appel à la transmission d'ultrasons par une pluralité de transducteurs orientés de manière à induire une cavitation dans une région choisie du cerveau au moins. Une source d'excitation permet d'actionner au moins les transducteurs choisis, à des phases qui diffèrent d'un transducteur à l'autre, par exemple pour compenser les déphasages (ou distorsions de phase) causés par le crâne sur l'émission d'ultrasons de chaque transducteur. Les ondes ultrason provenant des transducteurs atteignent ainsi la région choisie en étant sensiblement en phase les unes avec les autres, par exemple à moins de 90° et, de préférence, à 45° et mieux encore à 20° les une de l'autre.
PCT/US1997/014760 1996-08-21 1997-08-21 Methodes et dispositifs pour administrer un traitement par ultrasons non invasif au cerveau a travers une boite cranienne intacte WO1998007373A1 (fr)

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US6612988B2 (en) * 2000-08-29 2003-09-02 Brigham And Women's Hospital, Inc. Ultrasound therapy
US6770031B2 (en) 2000-12-15 2004-08-03 Brigham And Women's Hospital, Inc. Ultrasound therapy
CA2543725A1 (fr) 2002-10-28 2004-05-06 John Perrier Dispositif medical a ultrasons
US7909782B2 (en) 2002-10-28 2011-03-22 John Perrier Ultrasonic medical device
US20090230823A1 (en) * 2008-03-13 2009-09-17 Leonid Kushculey Operation of patterned ultrasonic transducers
CA2725453C (fr) 2008-04-30 2021-09-14 Milux Holding S.A. Stimulation cerebrale
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US9116098B2 (en) 2013-02-12 2015-08-25 General Electric Company Ultrasonic detection method and system
US9482645B2 (en) 2013-05-17 2016-11-01 General Electric Company Ultrasonic detection method and ultrasonic analysis method
US10589129B2 (en) * 2016-09-14 2020-03-17 Insightec, Ltd. Therapeutic ultrasound with reduced interference from microbubbles
US20220023668A1 (en) * 2020-07-21 2022-01-27 Liminal Sciences, Inc. Ultrasound annular array device for neuromodulation

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US5431621A (en) * 1984-11-26 1995-07-11 Edap International Process and device of an anatomic anomaly by means of elastic waves, with tracking of the target and automatic triggering of the shootings
US5158071A (en) * 1988-07-01 1992-10-27 Hitachi, Ltd. Ultrasonic apparatus for therapeutical use
US5316000A (en) * 1991-03-05 1994-05-31 Technomed International (Societe Anonyme) Use of at least one composite piezoelectric transducer in the manufacture of an ultrasonic therapy apparatus for applying therapy, in a body zone, in particular to concretions, to tissue, or to bones, of a living being and method of ultrasonic therapy
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