WO2011092683A1

WO2011092683A1 - Non-invasive ultrasound treatment of subcostal lesions

Info

Publication number: WO2011092683A1
Application number: PCT/IL2011/000063
Authority: WO
Inventors: Benjamin Sabbah; Igal Ilouz; Anat Roytberg
Original assignee: Livesonics Ltd.
Priority date: 2010-02-01
Filing date: 2011-01-20
Publication date: 2011-08-04

Abstract

High Intensity Focused Ultrasound probe for noninvasive therapeutic treatment of subcostal lesions. A first array of ultrasound transducers produce a first ultrasound beam having at least one frequency f2 and focus the beam at a target volume in the lesion and a second array of ultrasound transducers produce a second beam having at least one frequency f1, which is less then f2, and focused at the same target volume such that interference between the first and second beams creates in the target volume an asymmetric ultrasound field having a larger negative pressure amplitude than positive one. The dimensions of said probe are such that most of the energy of the first beam passes through interspaces between ribs. A method of using the probe for non-invasive therapeutic treatment of subcostal lesions comprising a cavitation stage in which bubbles are generated in the target volume followed by a heating stage.

Description

NON-INVASIVE ULTRASOUND TREATMENT OF SUBCOSTAL

LESIONS

Field of the Invention

The present invention relates to the field of ultrasound (US) technology. Specifically the invention relates to a system for creating High Intensity Focused ultrasound (HIFU) waves for medical applications at tissues located within the rib-cage.

Background of the Invention

The use of US for medical diagnostics and therapy is well known. Diagnostic techniques are based on the production and transmission of low energy US waves into the body, and detection of the scattered echoes from the scanned region. HIFU therapeutic methods are generally based on the use of focused ultrasonic energy to provide high intensity pressure waves at specific locations. The focused high intensity pressure waves lead to localized cavitation, heating or mechanical pressure, or combinations of these effects, that may be utilized for therapeutic purposes, by ablation, by causing mechanical damage, neovascularization, etc., within the medical target.

Rapid hyperthermia resulting in tissue ablation or necrosis has been proven to be a useful therapeutic modality, for example, in destroying cancerous tissue. Such applications are based on the use of high intensity focused beams of ultrasonic waves produced by either a single high power transducer or by focusing the energy from several transducers at the same location. Clinical application of HIFU is foreseen as the future non-invasive modality of choice, replacing other minimal invasive methods, e.g. RF/Laser ablation. The ability to focus the US energy at the specific location, with minimal absorption by neighboring tissue, by the tissue in between the transducer and that location, or by the tissue beyond that location is the limiting factor for implementation of this technology for therapeutic purposes.

In body fluids such as blood or in the intercellular fluids in living tissue, the application of ultrasonic energy often leads to the creation of bubbles which grow in volume by a process known as rectified diffusion. Conventional US signals are generated by transducers powered by a sinusoidal electrical voltage waveform that is applied to the transducer. When the applied electrical signal is in the frequency range of the transducer's frequency bandwidth and the signal is at least a few cycles long, the pressure wave generated by the transducer is of similar shape. The waves that are emitted from the transducer travel through the media as longitudinal waves (the transverse waves usually attenuate very rapidly and thus can be practically ignored) having alternating compression and decompression regions corresponding to the positive and negative portions of the waveform. When the pressure wave passes through a fluid, gases trapped inside dust motes or other particles in the fluid will be drawn out from the fluid forming a small bubble. If the acoustic power density is small, then the bubble will oscillate around a relatively constant radius. This process is known as stable cavitation. If the power density is high enough, then gas diffuses into the bubble during the de-compressing half-cycle of the ultrasonic wave and diffuses out from the bubble during the compression half-cycle. The rate of the diffusion is proportional to the radius of the bubble and therefore the rate of diffusion into the bubble, which occurs when the bubble has expanded during the de-compressing phase, exceeds that of the rate of diffusion out of the bubble, which occurs when the bubble has been compressed during the compressing phase. The net result is that the radius of the bubble increase as the bubble oscillates. This process is known as rectified diffusion. If this process is continued, the radius of the bubble reaches a critical value where it can be no longer remain stable and the pressure caused by the next compression phase will cause the bubble to implode. The critical value depends on the power and the frequency of the ultrasonic energy.

At therapeutic intensities, the hyperthermia caused by the focused US energy is often accompanied by bubble activity. In vitro and in vivo experiments alike have shown that under certain conditions bubble activity can give rise to a doubling of the heating rate (MRI-guided gas bubble enhanced ultrasound heating in in-vivo rabbit thigh, Sokka S.D., King R., Hynynen K., Phys. Med. Biol. 48, pp. 238-241, 2003).Others have used cavitation suppression techniques to reduce the micro-bubbles formation and make accurate ablation (Kentaro Tasaki, Takehide Asano, Kazuo Watanabe, Hiroshi Yamamoto, Division of Surgery, Chiba Cancer Center, Chiba, Japan. In "Therapeutic Ultrasound", the proceeding of the 2^nd International Symposium on Therapeutic Ultrasound, ISTU 2002). But more precisely, it was reported that a several times larger volume of tissue was coagulated with HIFU in a canine prostate with administration of Albunex Contrast Agent (CA) than without it. Kidneys were exposed to HIFU at 3.2 MHz in degassed saline, and when CA Optison was administrated at the dose of 0.2 ml/Kg, the temperature elevation induced by HIFU exposure was more than doubled. The effect that tissue ultrasonic absorption will be doubled when contrast agent is added to the tissue at -0.1 Billion microbubbles/Kg was theoretically predicted (by the study of S Umemura, K Kawabata, K Sasaki, published at "Enhancement of sonodynamic tissue damage production by second-harmonic superposition: Theoretical analysis and its mechanisms", IEEE-TUFFC, Vol. 43, pp. 1054- 1062, 1996). It was also experimentally verified (by R. Glynn Holt, Ronald A. Roy, "Measurements of bubble-enhanced heating from focused MHz frequency ultrasound in a tissue -mimicking material", Ultrasound in Medicine & Biology, Vol. 27, Issue 10. pp.1399-1412, 2001).

The fact that the presence of the bubbles would have a positive influence on the thermal treatment is somewhat unexpected in view of earlier efforts to separate the cavitation and heating effects, since the micro-bubbles shadowed the region to be treated and might endanger the patient (Bailey MR, Couret LN, Sapozhnikov OA, Khokhlova VA, Ter Haar G, Vaezy S, Shi X, Martin R, and Crum LA, "Use of overpressure to assess the role of bubbles in focused ultrasound lesion shape in vitro, Ultrasound Med Biol, 27 (5). 695-708, 2001). In US 5,601,526, for example, it is described how the presence of cavitation can lead to unwanted tissue destruction both inside of and outside of the focal zone, when applying focused acoustic power to achieve thermal effects. This patent teaches a method and apparatus for applying thermal and cavitation waves either separately or together so as to reduce the interaction between them. US 5,573,497 teaches a method and apparatus for supplying the maximum thermal energy to a tissue while reducing or preventing cavitation. Despite the effort made in the past to separate the thermal from the cavitation effects of HIFU on living tissue, the above cited publications attest to the difficulty and importance of harnessing the energy concentrating effects of bubbles to do useful clinical work when exposed to ultrasound. The dominant heating mechanism depends on bubble size, medium shear viscosity and frequency-dependent acoustic attenuation. The bubble size distribution, in turn, depends on insonation control parameters, e.g acoustic pressure and pulse duration; on medium properties, notably the concentration of the dissolved gas; and on bubble shape irregularities that facilitate instabilities. The general heating problems in HIFU treatment of lesions that are not obscured by ribs will now be discussed. For purposes of this discussion some useful definitions are:

1. Acoustic Gain:

The acoustic gain is the ratio of ultrasonic field amplitude at the focus to that at the transducer surface. The acoustic gain GA depends on: f ₌ tatf? focus ^transducer

A ~ AmV tw eA n_M. ~ Z X t ¹⁰ (equation 1)

Where:

• Atransducer is the area of the transducer surface

• f is the ultrasonic frequency

· Z is the distance to the target

• C is the speed of sound.

2. Temperature Gain:

The temperature gain is the ratio of the temperature rise at the focus, ATfocus to that at the skin surface, AT_skin, which is in contact with the transducer surface. The temperature gain G_AT is given by: r (equation 2)

Where:

· a(f)focus and (f)_skin are the absorption coefficients of the tissue type at the focus and at the skin respectively.

• e ^a®^AtZ is the attenuation of the US beam along the beam from the probe to the target. a(f At is attenuation coefficient of the US beam in the body.

· tfocus and t_skin are the time that each of the regions is exposed to the ultrasonic fields. 3. Number of treated targets:

The number of target zones that need to be treated in order to ablate the entire lesion depends on the ratio of the entire volume of the lesion to that of the effective focal zone. This also determines the ratio of t_skin to tfocus^'

_. ^' _. tskin (equation 3) targets ~~ . ^'— _*£

The objective of HIFU treatment of lesions is to heat the entire lesion to >60° C with minimal heating of surrounding tissue, e.g. the skin, in order to ablate the entire lesion without damaging any surrounding tissue. In order to optimize the efficiency of HIFU treatment and based on the three definitions above, the function of the temperature gain G_AT that should be maximized is: focus on 4)

Z X C ^t rgets ^ ^*) skin

From this function the two major problems with the method can easily be seen:

Firstly the transducer is constantly in close proximity to a specific area of skin, while the lesion volume is continuously scanned one element after the other, by the relatively much smaller volume of the effective treatment focal volume. Thus "time at the rib/skin" is much larger than "time at a target point". Therefore for a large lesion the result will be considerable heating of the skin, as Ntargets becomes large.

Secondly, the attenuation along the beam _e ^~ ^^AtZ, reduces the energy provided to the target and becomes a problem. The general conditions are a(f)f_0Cus /a(f)_skin ~1. Therefore, the maximum value of G_AT is achieved by maximizing

(^A*™**»r- ^Xff X e- mZ (equation s)

\ z xc ■

In all conventional solutions to the above problems the transducer is selected to be the largest size practically feasible, in order to increase

Atransducer.

The optimal frequency as a function of the transducer (skin) to lesion distance Z, for a typical absorbing medium of 0.7dB/cm/MHz, is shown in Fig. 1. This figure explains why all conventional methods of HIFU treatment of lesions operate in the frequency range of ~ 2 MHz.

To date, US techniques have been applied therapeutically to varied medical conditions such as treatment of benign and malignant tumors of many types, e.g. in the prostate; varicose effects; destruction of gall and kidney stones; treatment of the esophagus; and unblocking blood vessels. Another important medical condition that is an important candidate for treatment with HIFU is the treatment of subcostal lesions, i.e. lesions that are obscured by ribs.

Worldwide, more than 800,000 new cases of liver cancer, either primary or metastatic liver cancer, which is still confined to the liver at the time of the diagnosis, are diagnosed each year. Primary and secondary liver malignancies are the most common types of tumors representing a significant source of morbidity and mortality worldwide. Surgical resection is considered the optimal definitive therapy, but only less than 20% of patients with either primary or metastatic liver disease are amenable to surgical therapy. Technical reasons, general patient condition or hepatic dysfunction even reduce the number of potential surgical patients (Steele G Jr. Natural history studies and the evolution of regional treatment modalities for patients with isolated liver metastases from primary colon and rectum carcinoma. Cancer Control 1996; 3:34-41). The high prevalence of liver malignancy together with the low number of surgical resection candidates has increased the interest in minimally invasive or non-invasive therapies as an alternative for primary and secondary unresectable malignant tumors and a complement for surgical resections.

Chemotherapy and nuclear radiation therapy, two common non-invasive therapies, are poorly effective treatment methods. Radio-Frequency Ablation (RFA), a minimal-invasive thermal treatment, offers a partial solution but provides very limited results. Despite this, since presently there is no better solution, RFA is the most widely used treatment as an alternative to surgical procedure for treatment of liver malignancies.

As described previously, the most advanced non-invasive technique, HIFU, produces thermal tumor ablation therapy, the aim of which is to destroy the entire tumor by using cytotoxic heat with minimal damage of adjacent normal tissue and sparing adjacent vital structures. The goal when using heating techniques is to raise the temperature of the tissue to be destroyed to 50°C to 100°C in order to produce coagulative necrosis while at the same time avoiding charring and vaporization of tissue.

Most liver malignancies and a lot of kidney malignancies are located behind the ribs. The location of the malignancies results in two main technical problems that prevent the use of HIFU as the ultimate efficient noninvasive therapeutic treatment of subcostal malignancies:

1. The objective of the HIFU treatment is to heat the entire lesion to >60° C with minimal heating of the ribs (as well as the skin and underlying tissue that would be exposed to the therapeutic US energy), in order to ablate the entire lesion without damaging any surrounding tissue. In order to optimize the efficiency of HIFU treatment of a subcostal lesion, the function of the temperature gain G_AT (equation 4) should be maximized.

In the specific case of subcostal lesions there are two specific significant problems with the maximization of G_AT, in addition to those related to the general HIFU problems (as described above):

· Firstly, the ribs have a significantly higher absorption coefficient than the target lesion. The ratio of the absorption coefficients a(f)fbcu_S/a(f)ribs is between 1/20 to 1/50 (Ultrasonic attenuation and velocity in bone, J A Evans and M B Tavakoli, Phys. Med. Biol., 1990, Vol. 35, No 10, 1387-1396), rather than the ratio a(f)fbcus/a(f)_skin ~1 in the regular (no ribs) case.

• Secondly, the ribs obscure and distort the ultrasonic beam, thus significantly reducing the acoustic gain GA by a factor of 2-5.

The combination of these two factors implies deterioration of the temperature gain G_AT by at least 2 orders of magnitude in the subcostal case relative to the regular (no ribs) case.

2. Ablation treatment of a malignancy must be done by using guidance means in order to ensure directing the treatment to the desired target and preventing damage to the surrounding tissue and by using control means in order to verify that the desired process is completed. Providing such means to non-invasive treatment such as HIFU is very difficult.

Two presently available methods that have been developed to provide a solution to the problems of applying HIFU to subcostal malignancies are known to the inventors: — The first solution is that of InSightec Ltd., Tirat Carmel, Israel. This method uses a complex US transducer consisting of thousands of phased array elements in order to:

a. avoid the ribs and reduce skin and underlying tissue from heating by switching off the US transmission from the elements facing the ribs;

b. achieve a small enough spatial US focal point by using only the working part out of the whole array; and

c. increasing the focusing gain.

The significant shortcomings of this solution are that the US system is very expensive. Also real-time MRI thermal imaging of the whole body volume which is being exposed to US is required for guiding and controlling the treatment. MRI is a very expensive and not an available tool especially for real time control during the whole treatment procedure

— Another solution is provided by Chongqing Haifu Technology (Chongqing, China). The method comprises a procedure consisting of surgical removal of the disturbing ribs followed by HIFU treatment of the "subcostal" lesions (without ribs). This solution is a pseudo non- invasive treatment that has significant shortcomings from the patient's point of view.

All existing available solutions to the problems that prevent the use of HIFU as the ultimate non-invasive treatment of subcostal malignancies are unsatisfactory from various aspects (commercial/user/patient). Therefore, although RFA offers only a partial solution and provides very limited results, it is currently still the most widely used treatment of primary and metastatic hepatic tumors. It is therefore a purpose of the present invention to provide an effective solution to the problems of applying HIFU for the treatment of subcostal lesions.

Further purposes and advantages of this invention will appear as the description proceeds.

Summary of the Invention

In a first aspect the invention is a method for utilizing High Intensity Focused Ultrasound (HIFU) for non-invasive therapeutic treatment of subcostal lesions. The method comprises two stages:

— A cavitation stage, wherein two or more ultrasound beams having different frequencies are focused at the same target volume located in the lesion such that interference between the beams creates in the target volume an asymmetric ultrasound field having larger negative pressure amplitude than the positive one. In this stage bubbles are generated in the target volume.

— A heating stage, which follows immediately the cavitation stage, wherein an ultrasound beam having frequency of 400KHz or less is focused at the target volume and wherein temperature in the target volume is elevated during the heating stage.

Each stage of the method is activated for no longer than several milliseconds and a cycle comprised of the two stages is repeated hundreds or thousands of times over a period of at least a second.

In embodiments of the method the diameter of the ultrasound beam having the highest frequency in the cavitation stage is such that most of the beam energy passes through the interspaces between the ribs. In embodiments of the method the lowest frequency of the ultrasound beam used in the cavitation stage is between 200KHz and 400KHz. In embodiments of the method the highest frequency of the ultrasound beam used in the cavitation stage is double the lowest frequency. Some or all of the ultrasound beams can have a sinusoidal waveform.

In embodiments of the method the lowest frequency ultrasound beam, is composed of two close frequencies f ± Αΐ, wherein Αΐ equals several KHz to several tens of KHz. The two close frequencies can be heterodyned by the bubbles to oscillate at a frequency equal to Δ

The bubbles generated in the target volume during the cavitation stage can be detected by a diagnostic ultrasound system for real-time ultrasonic guidance, monitoring and control of the treatment. Embodiments of the method comprise interrogating the region containing the oscillating bubbles by a Doppler detection system, using Doppler signals generated according to the oscillating frequency Δί. The Ultrasound system can be operated in one or more of the following modes: color Doppler mode, Doppler CW mode, and Doppler pulse mode.

In embodiments of the method wherein the diameter of the ultrasound beam having the highest frequency in the cavitation stage is such that most of the beam energy passes through the interspaces between the ribs, the highest frequency in the cavitation stage can additionally be used to detect ultrasound echoes scattered by the bubbles to monitor the density and location of the bubbles.

In embodiments of the method of the first aspect of the invention an ultrasound beam having a frequency higher than 2MHz is focused at a target located in the lesion to detect ultrasound echoes scattered by bubbles in order to monitor the density and location of the bubbles wherein most of the energy of the beam passes through the interspaces between the ribs.

In embodiments of the method of the first aspect of the invention a 2D/3D diagnostic ultrasound system can be used to image the lesion. In these embodiments a position sensing system is employed for fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the 2D or 3D ultrasound images made with the 2D/3D diagnostic ultrasound system and with the focal zone monitoring, which are taken in real-time during the HIFU treatment.

In a second aspect the invention is a High Intensity Focused Ultrasound (HIFU) probe for non-invasive therapeutic treatment of subcostal lesions. The probe comprises:

— a first array of ultrasound transducers having at least one transducer designed to produce at least one first ultrasound beam having at least one frequency Ϊ2 and to focus the beam at a target volume located in the lesion;

— a second array of ultrasound transducers having at least one transducer designed to produce at least one ultrasound second beam having at least one frequency fi, which is less then Ϊ2, and focused at the same target volume located in the lesion such that interference between the first and second beams creates in the target volume an asymmetric ultrasound field having a larger negative pressure amplitude than the positive one.

The dimensions of said probe are determined such that most of the energy of the first beam passes through the interspaces between the ribs. In embodiments of the probe the first ultrasound transducer array is an inner transducer array and the second transducer array is an outer transducer array. The second transducer array can have an annular cross section. In embodiments of the probe the diameter of the second ultrasound transducer array is between 10cm and 20cm. In embodiments of the probe the first ultrasound transducer array has a circular cross section having diameter between 1.5cm and 5cm. The inner ultrasound transducer array has an elliptical cross section whose maximum dimensions are: long axis between 2cm and 6 cm and short axis between 1.5cm and 5cm.

In embodiments of the probe ft is between 200KHz and 400KHz. In embodiments of the probe Ϊ2 is double ft. Some or all of the ultrasound beams can have a sinusoidal waveform. In embodiments of the probe ft is split into two close frequencies fi ± Αΐ, wherein Αΐ equals several KHz to several tens of KHz. In embodiments of the probe Αΐ equals between 2KHz and 25KHz.

In embodiments of the probe the second transducer array is comprised of several separate sub-elements that are operated at the same ultrasound frequency f_1; but are driven separately with different electronic signal phases. In embodiments of the probe the second transducer array is divided into two sections in order to enable spatial division of the ultrasound beam having frequency fi into two separate beams with different frequencies ΐι±Αΐ. In embodiments of the probe wherein the second transducer array is comprised separate sub-elements, the sub-elements are divided into two sections in order to enable spatial division of the ultrasound beam having frequency fi into two separate beams with different frequencies ΐι±Αΐ.

The first ultrasound transducer array that operates at frequency Ϊ2 and which is already focused on the treated target volume can be adapted to enable it to be operated in the Doppler CW or Pulse mode as a sensor probe for detection of the ultrasound echoes scattered by the bubbles. The probe of the invention may comprise an additional ultrasound transducer designed to produce an ultrasound beam having a frequency higher than 2MHz and adapted to enable it to act as a Doppler detector element. In this case the additional transducer is designed to be focused on the same focal point as the other ultrasound transducers of the probe and is embedded within the probe in a position where it can be aligned in front of the interspaces between the ribs, in order to enable transmission and detection of the high frequency beam only through the inter-ribs space. In a third aspect the invention is a system for non-invasive therapeutic treatment of subcostal lesions. The system of this aspect comprises:

— a High Intensity Focused Ultrasound (HIFU) probe according to the second aspect of the invention; and

— a diagnostic ultrasound system, which is adapted to be used for real- time ultrasonic guidance, monitoring and control of the treatment.

In a fourth aspect the invention is a system for non-invasive therapeutic treatment of subcostal lesions. The system of this aspect comprises:

— a High Intensity Focused Ultrasound (HIFU) probe according to the second aspect of the invention, wherein either the first ultrasound transducer array that operates at frequency Ϊ2 and which is already focused on the treated target volume is adapted to enable it to be operated in the Doppler CW or Pulse mode as a sensor probe for detection of the ultrasound echoes scattered by the bubbles or said probe comprises an additional ultrasound transducer designed to produce an ultrasound beam having a frequency higher than 2MHz and adapted to enable it to act as a Doppler detector element that focuses on the same focal point as the other ultrasound transducers of the probe and is embedded within the probe in a position where it can be aligned in front of the interspaces between the ribs, in order to enable transmission and detection of the high frequency beam only through the inter-ribs space; and

— a position sensing system adapted to enable fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the focal zone monitoring, which is taken in real-time during the HIFU treatment.

In a fifth aspect the invention is a system for non-invasive therapeutic treatment of subcostal lesions. The system of this aspect comprises:

— a diagnostic ultrasound system, which is adapted to be used to image the lesion.

The system of the fifth aspect of the invention may additionally comprise a position sensing system adapted to enable fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the 2D or 3D ultrasound images made with the 2D/3D diagnostic ultrasound system and with the focal zone monitoring, which are taken in real-time during the HIFU treatment.

All the above and other characteristics and advantages of the invention will be further understood through the following illustrative and non-limitative description of preferred embodiments thereof, with reference to the appended drawings.

Brief Description of the Drawings

— Fig. 1 is a graph showing the optimal US frequency as a function of the transducer (skin) to lesion distance Z, for a typical absorbing soft tissue medium;

— Fig. 2 and Fig. 3 are graphs of temperature vs. time that show typical experimental results of the heating using 350 KHz US without (Fig. 2) and with (Fig. 3) bubble formation at the focus;

— Fig. 4 and Fig. 5 show experimental results obtained that demonstrate the efficiency of the heterodyne method;

— Fig. 6(a) to Fig.6(d) schematically show front views of different embodiments of the focused HIFU probe of the invention;

— Fig. 7 shows the calculated asymmetric US field resulting from superposition of

and f₂=700KHz;

— Fig. 8 shows the measured asymmetric US field resulting from experimental superposition of fi=350KHz and f2=700KHz;

— Fig. 9 shows experimental results of an ex-vivo experiment carried out using a meat sample located behind the ribs of a pig using the cavitation and heating stages ;

— Fig. 10 shows experimental results of an ex-vivo experiment carried out using a meat sample located behind the ribs of a pig using only the heating stage of the method of the invention; — Fig. 11 symbolically shows a front view of an embodiment of the focused HIFU probe of the invention that comprises an embedded Doppler detector element.

— Fig 12 - schematically illustrates a front view example for the most general embodiment of the probe of this invention.

Detailed Description of Embodiments of the Invention

The present invention is a system and a method that utilizes HIFU for noninvasive therapeutic treatment of subcostal lesions. Relatively low frequencies are used and focused to generate microbubbles at the target, to produce stable or transient cavitation and localized heating. Additionally, the system and method provide treatment guidance and monitoring by employment of diagnostic US measurements or imaging. The method of the invention is a two stage process. The first stage comprises a cavitation process for highly efficient and localized generation of microbubbles from dissolved gases within the intra-cell liquid/blood in a desired target volume. This stage selectively increases the potential ratio between US energy absorption at the target volume at the lesion site and other surrounding tissue in the body including the ribs (bone).

The process used in the first phase for a highly efficient and localized generation of microbubbles in the target volume is based on the following:

1. When liquid contains dissolved gas, microbubbles can be generated by a high negative pressure. In intra-cellular liquid/blood within the body this requires a pressure of a few mega-Pascals (Mp's) that can be achieved by applying US in the order of > KW/cm², with the exact value depending on the US frequency.

2. When the pressure P is negative the direction of gas diffusion is mostly from the liquid into the bubbles. 3. When the pressure P is positive the direction of gas diffusion is mostly from the bubbles into the liquid.

4. Gas diffusion is proportional to ΔΡ*4πΙ ², where R is the bubble radius.

Symmetric US fields, i.e. fields having equal negative and positive pressure amplitudes, cause rectified diffusion, i.e. the size of the bubbles increases with time, since during the duration of the negative pressure the diameter of the bubbles is larger than during the duration of the positive pressure, thus more gas defuses into the bubble than defuses out of the bubble into the liquid.

5. An asymmetric US field, with a larger negative pressure amplitude than the positive one, i.e. peak | P. | > peak I P+ 1 , generates microbubbles more efficiently than a symmetric US field by the factor P./P+ [Church C. A theoretical study of cavitation generated by an extra corporal shock wave lithotripter. J. Acoust. Soc. Am. 1989; 86: 215-227]. This means that the US energy required to generate microbubbles is reduced by a factor of (P./P+)². Thus using an asymmetric US field, with larger negative pressure amplitude than positive pressure amplitude enables generation of microbubbles without undesired thermal heating of the target and allows using lower power sources.

6. An asymmetric US field can be applied by the interference of two or more US beams having different frequencies.

7. If all of the beams having different US frequencies are focused at the same focal point and their phases are aligned properly, then the interference that produces the US field with the desired P./P+ ratio is highly localized in terms of bubbles generation to a very small effective lateral spot size and small axial focus depth. This is in comparison with the spot size and focus depth due to each individual US frequency with the same "Numerical Aperture" (which are the parameters that define the spot size and focus depth). The local cavitation (bubble generation) at the target volume is followed immediately by a second stage of the process which comprises focusing low frequency (200-400 KHz) at the target volume. In this relatively low US frequency regime of 200-400 KHz the US is not significantly absorbed by the ribs or by other soft tissue, e.g. skin or the liver itself. Also at these low US frequencies the bubbles formed in the localized target volume have a significantly higher US absorption coefficient. Thus it has been found in experiments carried out by the inventors that a(f)f_0Cus/a(f)ribs is increased to be -10, compared to the situation in the regular case where a(f)f₀cus/a(f)ribs is at most 1/20. Referring to equation 4 it is seen that the temperature gain G_AT is increased locally in the target by the same ratio.

Thus the presence of microbubbles only in the target volume provides a very selective increase in the temperature of the target tissue. The efficiency of the two stage method of the invention was experimentally verified using a Gel phantom. Fig. 2 and Fig. 3 are graphs of temperature (measured at the focus and at 5mm above and below the focus) vs. time. The graph in Fig. 2 shows typical experimental results of the heating using only the 350 KHz US frequency. Fig. 3 shows the heating after bubble formation at the focus using the interference of US beams at 350KHz and 700KHz to generate an asymmetric US field, for highly efficient and localized generation of microbubbles.

In embodiments of the invention the low frequency, which is used for the second (heating) stage, is composed of two close frequencies that differ by several KHz to several tens of KHz. These two frequencies are heterodyned by the bubbles to oscillate at a frequency equal to the difference between those two frequencies. The microbubbles have significant response to such oscillating frequency; since it is on the same order of magnitude as their typical resonance, which depends on their average size, and therefore their size is effectively enlarged. The US absorption coefficient increases as the bubble size increases. Therefore, splitting of the low frequency, which is used for the heating stage, into two close frequencies enables the heterodyning effect to be utilized in order to maintain and even enhance the US absorption contrast, i.e. the factor a(f)f₀cus/a(f)ribs by maintaining the presence of relatively large microbubbles during the heating stage that follows the stage of microbubble generation. The hetrodyne method enables the selective heating process to be improved by reducing the US power that is required to increase the temperature at the target and thus to reduce heating of the ribs and all other surrounding tissue. Fig. 4 and Fig. 5 show experimental results obtained using an ex-vivo meat sample that demonstrates the efficiency of the heterodyne method. Fig. 4 and Fig. 5 are graphs of temperature (measured at the focus and at 5mm above and below the focus) vs. time showing typical experimental results of the heating using 350 KHz US at 100% power without heterodyning (Fig. 4) and using 345KHz and 355KHz with heterodyning at 40% power (Fig. 5) both after bubble formation at the focus.

The timing of the two stages of the method of the invention must be optimized to maximize the rate of the heat generation. The process of generating microbubbles becomes saturated after several milliseconds. Also the efficiency of the selective heating process decreases after several milliseconds due to the decrease with time of the density of microbubbles in the target. Therefore, the optimized timing requires activation of each stage for no longer than several milliseconds.

The basic cycle of the two stages of the basic process takes place in a time scale of milliseconds. The basic cycle is repeated hundreds or thousands time over a period of several seconds, in order to produce accumulation of the desired heat rise in the target being treated. The graphs in Fig. 4 and Fig.5, which have time scale of seconds, show the full heating process, which comprises a few thousand cycles of the two stage basic process. Each sonication section of the cycle is comprised of a "pulse mode" lasting on the order of milliseconds during which two crystal elements and frequencies are excited simultaneously in order to generate the microbubbles followed immediately by a "continuous mode" also lasting on the order of milliseconds during which only the low frequency is excited for the selective heating process. These two modes are repeated one immediately after the other without pause for a few seconds. The acoustical power transmitted by the transducer is on the order of tens to hundreds of Watts. US frequencies in the range of 200-400 KHz are used for the second (selective heating) stage of the process. The use of these frequencies, which are considerably lower than the frequencies conventionally used in HIFU (see Fig. 1) decreases the effect of beam defocusing by the ribs.

Bubbles can be seen by diagnostic US means. Therefore, generation of bubbles in the target being treated enables conventional commercial diagnostic US systems to be used for real-time ultrasonic guidance, monitoring and control of the treatment as will be discussed herein below.

A special probe has been invented, wherein the transducer design is optimized to implement the method of the present invention, taking into account the specific problems discussed herein that are related to HIFU treatment of subcostal lesions.

Fig. 6(a) schematically shows a front view of one embodiment of the focused probe of the invention. The probe comprises two elements. The inner element, labeled H in Fig. 6(a), is a transducer having circular cross section that is designed to produce the high frequency US. The second element, labeled L in Fig. 6(a), is a transducer having an annular cross section that is designed to produce the low frequency US. The active area of the probe is spherical to focus the US energy emitted by the two transducers.

The outer annular element has the largest size that is practically feasible, typically in the range of 10-20 cm, in order to increase Atransducer and thus maximize GA (see equation 1). The resonance frequency fi of this element is selected to be in the range of 200-400 KHz since US of this frequency is not significantly absorbed by the ribs or by the skin tissue. The resonance frequency Ϊ2 of the inner transducer element is selected to be and therefore in the range of 400-800KHz. The chosen frequency ratio is f2=2fi in order to ensure the periodicity (relative to the lower frequency) of the asymmetric waveform having a larger negative pressure amplitude than positive one of the US field that is produced in the focal region by interference of these two frequencies.

At the higher US frequency regime f2 the US absorption coefficient in the ribs, i.e. a(f)_ribs, is significantly higher than for fi. Therefore, the size and shape of the central element is designed to enable transmission of the higher frequency only through the interspaces between the ribs. Typically, the diameter of the circular inner element in the embodiment shown in Fig. 6(a) is in the range of 1.5-5 cm, depending on the radius of curvature of the focused probe, the working distance between the probe interface and the focal point, and hence on the distance to the ribs.

Fig. 6(b) schematically shows a front view of another embodiment of the focused probe of the invention. In this embodiment the inner element has an elliptical cross-section. As in the embodiment shown in Fig. 6(a), the active area of the two elements array probe is spherical. In this embodiment the long axis of the inner elliptical element should be aligned with the direction of the inter-ribs space. A typical dimension for the length of the long axis of the elliptical element is ~5cm and for the length of the short axis in the range of 1.5-5 cm (note that as the short axis approaches 5cm, the embodiment in Fig. 6(b) becomes essentially equivalent to that of Fig. 6(a)).

The higher US transmission at frequency Ϊ2 from the central element, together with the lower US frequency fi, which is transmitted from the outer element, enables the creation of the desired asymmetric US field, with larger negative pressure amplitude than the positive one, thereby enabling highly efficient and localized generation of microbubbles.

In the first stage the outer transducer and the inner one are activated simultaneously for the purpose of generating microbubbles. Immediately following the microbubble generation stage the inner element is shut off and the outer transducer element transmits a waveform at a single low frequency fi (200-400 KHz) to cause the selective heating in the target. In the preferred embodiment of the invention, for the purpose of the heating process the low frequency fi is preferably split up to ΐι±Δΐ, where Αΐ is between 2KHz to 25KHz in order to utilizes the heterodyning effect on the microbubbles.

It is well known that a spherical focused US transducer has a focal zone depth that is proportional to the US wavelength and to the transducer f- number, which is the ratio between the focus and the aperture diameter [See for example: Fry W and Dunn F., "Ultrasound: analysis and experimental methods in biological research" in: Physical Techniques in Biological Research", Ed. Nastuk, W.L., Vol. IV: 261-394, Academic Press New York, 1962]. For the probe of the invention f₂=2fi and Όι>2Ό₂. Therefore, the focal zone depth of f2 is longer than that of fi, according to the ratio between the chosen designed diameters of these two transducer elements. This feature can be utilized in order to improve efficiency of the HIFU focal zone movement to multiple depths, as will be explained in the next paragraph.

Another embodiment of the focused probe of the invention is schematically shown in Fig. 6(c). In this embodiment of the probe, as in the previously described embodiments, the active area of the probe is spherical and the outer annular element, which operates at the low frequency fi, has the largest size that is practically feasible. In this embodiment the outer annular element is comprised of several separate annular sub-elements. Five sub-elements are shown in Fig. 6(c) but the actual number can be larger or smaller, e.g. from two to ten. These separate sub-elements are operated at the same US frequency f_1; but are driven separately with different electronic signal phases. Each of the sub-elements generates an US wave with an independent initial phase, such that the waves interfere constructively with each other at a location that ^" depends on the corresponding phases. If all of the initial phases of the US waves generated by each sub-element are the same, then this is a degenerate case resembling the one element spherical transducer shown in Fig. 6(a), where the focal point is in the geometrical center of the sphere. An US beam that generates a constructive interference (focal point) on the centric axis at a specific distance from the transducer, which is not necessarily located at the geometrical focal point of the spherical transducer, can be achieved by adjusting the phase of the signal that drives each sub-element of the outer annular element to the required focal distance [Bjoren AJ, Angelsen, "Ultrasonic Imaging", Emantec AS, Trondhein, Norway, 2000]. In this embodiment of the probe of the invention the focal zone of the fi beam may be focused at variable depths while it is still kept within the focal zone depth of the £2 beam. Therefore the conjoint focal zone can be moved to multiple depths within the body in order to improve the movement efficiency of the HIFU treated focal zone within the target. In order to optimize the use of the heterodyning effect on microbubbles only in the focal point, the outer annular element of the probe is divided into two sections as shown schematically in Fig. 6(d). This division is used to divide spatially the low frequency US beam at ΐι into two separate beams with different frequencies fiiAf. The heterodyning effect can only occur when these two beams interfere, i.e. in the conjoint focal point.

Fig. 6(d) shows an embodiment in which the outer element has been divided into two separate elements, marked L¹ and L². Before dividing it into two, the outer element can be a single element as in Fig. 6(a) or Fig. 6(b) or it can be an array of composed of multiple annular sub-elements as in Fig. 6(c). The inner element, marked H can be circular as in Fig. 6(a) or Fig. 6(c). The inner element can be non-circular, for example elliptical as in Fig. 6(b), only in case that the outer element is a single element and not an array. The active area of the probes according to this embodiment is spherical.

In this invention the superposition of fi and Ϊ2 at their joint foci produces the desired asymmetric US field, with larger negative pressure amplitude than the positive one. As said previously a US field having this property generates microbubbles more efficiently by the factor | P. | / 1 P+ 1 =2. This means that the US power required to generate microbubbles is reduced by a factor of (P-/P+)²=4 in comparison with standard symmetric sine wave US field. Fig. 7 shows the calculated asymmetric US field resulting from superposition of

and f₂=700KHz.

Fig. 8 shows the measured asymmetric US field resulting from experimental superposition of

and f₂=700KHz. Fig. 9 shows experimental results of an ex-vivo experiment carried out using a meat sample located behind the ribs of a pig. The temperature measurements were made using thermocouples located at the focus, 5mm above the focus, at the transducer surface, and at four representative locations on the ribs where a temperature rise caused by exposure to and absorption of US energy could be expected to be detected. The sonication was carried out using both the cavitation and heating stages cyclically each for several milliseconds and repeating the basic cycle for a total duration of 5 seconds. Note that the horizontal axis in Fig. 9 (and Fig. 10) shows time in seconds, i.e. the whole X-axis represents 52 seconds. The results show an efficient and selective temperature rise at the focus and essentially no temperature rise at any of the other locations.

Fig. 10 shows experimental results from an experiment carried out under exactly the same conditions as the previous one, but this time without operating the first stage of the process, i.e. microbubble generation. The results show that in this case no significant temperature elevation is achieved at the focus while the results for all the other thermocouples are the same as in the previous case. These two experiments demonstrate the dependence of the high localized and efficient temperature elevation achieved using this method on the microbubbles generation stage.

A method for improved detection and monitoring of bubble assisted HIFU heating of tissue, using a commercial diagnostic US system is now described. The method enables guiding and locating the focal treated zone into the target lesion, which is imaged by using a conventional two dimensional (2D) diagnostic US system or three dimensional (3D) US imaging. The method is based on Doppler detection and monitoring of the treated region.

The outer element of the HIFU transducer transmits the low frequency for the heating stage. This low frequency waveform is composed of 2 close frequencies, fi±Af, where Ai is between several KHz to several tens of KHz. These two frequencies are heterodyned by the bubbles to oscillate at a frequency equal to ϊ. When the region containing the oscillating bubbles is interrogated by a Doppler detection system, Doppler signals are generated according to the oscillating frequency Δ The intensity of the Doppler signals will be proportional to the intensity of the low US frequency transmission and to the bubble density. When using color Doppler in the variance mode, multiple colors will be generated. When Doppler CW mode is operated, much higher bubble modulation frequencies can be detected and therefore can also be allowed to be generated. When Doppler Pulse Mode is used, it is desirable to operate the Doppler system at a Pulse Repetition Frequency that will allow measurement of the bubble oscillation signal without aliasing. Thus, the bubble oscillation frequency can be varied according to depth, in order to allow non-aliased measurement. The central element of the HIFU transducer that operates at the higher frequency Ϊ2 and which is already focused on the same treated target can also be used as a sensor probe for detection of the US echoes scattered by the bubbles. This detection system is operated in the Doppler mode (CW/ Pulse mode) in order to improve the detection and monitoring of the bubble region. This enables target point monitoring of the process of generating the bubbles in the focal zone. The intensity of the detected signal depends on the bubble density. The focus point of the transducer is located in a defined position, which lies on the centric axis that is perpendicular to the transducer surface; but, there is an uncertainty regarding the exact actual distance between the transducer surface and the effective focal treated zone. The actual distance can be determined by measuring the time delay between US transmission and the detected echoes signal. On the one hand this is only a non-imaging detection system, but on the other hand it has a focus only in the targeted region of interest and therefore filters out undesired noise from the surrounding area. For improved spatial resolution monitoring of the generated bubbles, the preferred operation of the detection probe is at higher US frequency than the Ϊ2 range (400-800KHz). Therefore, another optional embodiment of the probe, which is shown schematically in Fig. 11, has been developed by the inventors. This embodiment of the probe includes an additional element that operates at a US frequency higher than 2MHz (hence, also higher than Ϊ2). This Doppler detector element is designed to be focused on the same focal point as the HIFU elements. In this embodiment of the probe, the Doppler detector element should be embedded within the probe in such a position where it could be aligned in front of the interspaces between the ribs, in order to enable transmission and detection of the high frequency (>2MHz) only through the inter-ribs space since, as described herein above, at this US frequency regime the US absorption coefficient in the ribs is significant. The probe that is shown schematically in Fig. 11 is only an example where the Doppler detector element is chosen to be embedded as the central element of the whole probe and is concentric with all other HIFU elements. When monitoring of the bubbles is operated simultaneously with the HIFU transmission mode that uses only the low frequency element (marked as L in the figures), the central detection element can be activated alone or may also be used together with the inner HIFU element (marked as H in the figures) while it can be used as a sensor. If both of these elements are used together for monitoring, two separate Doppler detectors provide detected data information for the monitoring of the bubbles.

The active area of all elements in the probe shown in Fig. 11 is spherical. The outer HIFU element, marked L (Low frequency), can (or can not) be divided into two separate elements, marked L¹ and L² as illustrated in Fig. 11. The outer area also can be an array separated into multiple annular sub- elements as shown in Fig. 6(c) or it can be any combination of both of these options. The inner HIFU element, marked H (High frequency), can be circular (as shown in Fig. 11) or non-circular (for example elliptical as shown in Fig. 6(b)). A non-circular inner element can be used only if the outer element is a single element and not an array. The HIFU element H can also be used as a sensor probe for bubbles monitoring; however, the central element (marked D in Fig. 11) can function only as a detector for monitoring of the bubbles.

The most general embodiment of the probe of the invention for utilizing HIFU for non-invasive therapeutic treatment of subcostal lesions might comprise larger number of US transducer arrays than those described in the previous examples. It might comprise some or all of the following:

• At least two HIFU transducer arrays, that transmit at leastTtwo different US waveforms, which are focused at the same location to produce interference that generates asymmetric US field, with larger negative pressure amplitude than the positive one, thereby creating the conditions that lead to highly efficient and localized generation of microbubbles from dissolved gases within the intra-cell liquid/blood in said target.

• At least one HIFU transducer array each with at least one element, designed to enable transmission of an US beam that most of its energy passes through the interspaces between the ribs, thereby do not heat the ribs.

• One or more HIFU transducer elements designed to enable transmission through the ribs of US beams of waveform with minimal absorption by the ribs.

• At least one transducer element, designed to enable monitoring of the generated bubbles by detection the US echoes scattered by the bubbles, as already described above. The elements of this detector are designed to be focused through the interspaces between the ribs on the same focal point as the HIFU elements. Fig. 12 schematically illustrates a front view example for the most general embodiment of the probe of this invention. The probe comprises three element arrays, labeled in Fig. 12 as L, H and D:

1. The outer HIFU array, each element labeled L in Fig. 12. Sixteen elements are shown in the schematic example illustrated in Fig. 12, although the actual number can be larger or smaller, but at least one. The frame shape of this array can have any cross section shape (circular, elliptical, polygon etc.). This outer array has the largest size that is practically feasible, with typical length dimension of side, axis or diameter of its frame (shape pending) in the range of 10-20 cm, that in order to increase Atransducer and thus maximize GA (see equation 1). This outer array is designed to transmit the low US frequencies with minimal absorption by the ribs although it pass through the ribs. The resonance frequency fi of each element in this array is selected to be in the range of 200-400 KHz, since US of this frequency is not significantly absorbed by the ribs or by the skin tissue. In order to optimize the use of the heterodyning effect on microbubbles only in the focal point, the outer array can be divided into two sections. This division is used to divide spatially the low frequency US beam at fi into two separate beams with different frequencies fi±Af. The heterodyning effect can only occur when these two beams interfere, i.e. in the conjoint focal point. The division into two sections is achieved by selecting part of the elements to be ΐι+Αΐ and part to be fi+Δί, not necessarily adjacent elements.

2. The inner HIFU array, each element labeled H in Fig. 12. Four elements are shown in the schematic example illustrated in Fig. 12, although the actual number can be larger or smaller, but at least one. The frame shape of this array can have any cross section shape (circular, elliptical, polygon etc.) and is designed to transmit the high US frequency beam in such a way that most of its portion can pass only through the interspaces between the ribs. The resonance frequency Ϊ2 of each element in the inner array is selected to be

and therefore in the range of 400-800KHz. 3. Optionally a detection array, for bubbles monitoring, which is embedded within the HIFU probe, each element labeled D in Fig. 12. Four elements are shown in the schematic example illustrated in Fig. 12, although the actual number can be larger or smaller, but at least one if this possibility is implemented within the HIFU probe. As described above, the elements of the inner HIFU array (labeled H) can also be used as a sensor probe for bubbles monitoring. In this case elements H and D are the same and are bi-functional. For improved spatial resolution monitoring of the generated bubbles, the preferred operation of the detection array is at higher US frequency than the Ϊ2 range (400-800KHz). In the embodiment that is illustrated in Fig. 12 the probe includes a separate detection array that operates at a US frequency higher than 2MHz (hence, also higher than Ϊ2. This detector array is designed to enable focusing on the same focal point as the two HIFU arrays and should be embedded within the probe in such a position where it could be aligned in front of the interspaces between the ribs, in order to enable transmission and detection of the high frequency (>2MHz) only through the inter-ribs space since, as described herein above, at this US frequency regime the US absorption coefficient in the ribs is significant. When monitoring of the bubbles is operated simultaneously with the HIFU transmission mode that uses only the low frequency array (marked as L in the figure), the detection array (marked as D in the figure) can be activated alone or may also be used together with the inner HIFU array (marked as H in the figure) if it is designed to be functional also as a sensor for bubbles monitoring. If both of arrays (D and H) are used together for this monitoring, actually two separate detectors provide simultaneously detected data information for the bubbles monitoring.

Each of the three transducer arrays might be operated by phased array technique in order to move the focus location within the desired target position without mechanical movement of the probe. Therefore in this case, the separate elements of each array are driven separately with different electronic signal phases. Each element generates an US wave with an independent initial phase, such that the waves interfere constructively with each other at a location that depends on the corresponding phases.

An independent commercial 2D/3D diagnostic US system is used in order to image the target lesion and the HIFU transducer is used to locate the position of the focal treated zone, as described above. The position of the 2D/3D US imaged lesion is taken relative to the diagnostic US probe while the position of the treated focal zone is taken relative to the HIFU transducer, which has a different position and orientation; therefore, an accurate 3D alignment between these two independent probes is required in order to guide the focal treated zone position into the target lesion. The required alignment can be achieved by using a commercial position sensing and tracking system, e.g. magnetic "Flock of Birds" (FOE) position sensing, or optical position sensing.

In another method for improved guiding and locating the focal HIFU treated zone into the target lesion, a commercial position sensing system (as described above) can also be employed for fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the 2D or 3D US images and with the focal point monitoring, which are taken in real-time during the HIFU treatment. Although embodiments of the invention have been described by way of illustration, it will be understood that the invention may be carried out with many variations, modifications, and adaptations, without exceeding the scope of the claims.

Claims

1. A method for utilizing High Intensity Focused Ultrasound (HIFU) for non-invasive therapeutic treatment of subcostal lesions, said method comprising two stages:

— A cavitation stage, wherein two or more ultrasound beams having different frequencies are focused at the same target volume located in said lesion such that interference between said beams creates in said target volume an asymmetric ultrasound field having a larger negative pressure amplitude than the positive one;

wherein bubbles are generated in said target volume during said cavitation stage; and

— a heating stage, which follows immediately said cavitation stage, wherein an ultrasound beam having frequency of 400KHz or less is focused at said target volume and wherein temperature in said target volume is elevated during said heating stage;

wherein, each stage is activated for no longer than several milliseconds and a cycle comprised of said two stages is repeated hundreds or thousands of times over a period of at least a second.

2. The method of claim 1, wherein the diameter of the ultrasound beam having the highest frequency in the cavitation stage is such that most of the beam energy passes through the interspaces between the ribs.

3. The method of claim 1, wherein the lowest frequency of the ultrasound beam used in said cavitation stage is between 200KHz and 400KHz.

4. The method of claim 1, wherein said highest frequency of the ultrasound beam used in said cavitation stage is double the lowest frequency.

5. The method of claim 1, wherein some or all of the ultrasound beams have a sinusoidal waveform.

6. The method of claim 1, wherein the lowest frequency ultrasound beam, is composed of two close frequencies f ± Αΐ, wherein Αΐ equals several KHz to several tens of KHz.

7. The method of claim 6, wherein the two close frequencies are being heterodyned by said bubbles to oscillate at a frequency equal to Αΐ.

8. The method of claim 1, wherein the bubbles generated in the target volume during the cavitation stage are detected by a diagnostic ultrasound system for real-time ultrasonic guidance, monitoring and control of the treatment.

9. The method of claim 8 comprising interrogating the region containing said oscillating bubbles by a Doppler detection system, using Doppler signals generated according to the oscillating frequency Af.

10. The method of claim 8, wherein the Ultrasound system can be operated in one or more of the following modes: color Doppler mode, Doppler CW mode, and Doppler pulse mode.

11. The method of claim 2, wherein the highest frequency in the cavitation stage is additionally used to detect ultrasound echoes scattered by the bubbles to monitor the density and location of said bubbles.

12. The method of claim 1, comprising focusing an ultrasound beam having a frequency higher than 2MHz at a target located in the lesion to detect ultrasound echoes scattered by bubbles in order to monitor the density and location of said bubbles wherein most of the energy of said beam passes through the interspaces between the ribs.

13. The method of either of claim 10 or claim 11, wherein a 2D/3D diagnostic ultrasound system is used to image the lesion.

14. The method of claim 13, wherein a position sensing system is employed for fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the 2D or 3D ultrasound images made with the 2D/3D diagnostic ultrasound system and with the focal zone monitoring, which are taken in real-time during said HIFU treatment.

15. A High Intensity Focused Ultrasound (HIFU) probe for non-invasive therapeutic treatment of subcostal lesions, said probe comprising:

— a first array of ultrasound transducers having at least one transducer designed to produce at least one first ultrasound beam having at least one frequency Ϊ2 and to focus said beam at a target volume located in said lesion;

— a second array of ultrasound transducers having at least one transducer designed to produce at least one ultrasound second beam having at least one frequency f_1; which is less then Ϊ2, and focused at the same target volume located in said lesion such that interference between said first and second beams creates in said target volume an asymmetric ultrasound field having a larger negative pressure amplitude than the positive one;

wherein, most of the energy of said first beam passes through the interspaces between the ribs.

16. A probe according to claim 15 wherein the first ultrasound transducer array is an inner transducer array and the second transducer array is an outer transducer array.

17. A probe according to claim 15 wherein the second transducer array has an annular cross section.

18. A probe according to claim 17, wherein the diameter of the second ultrasound transducer array is between 10cm and 20cm.

19. A probe according to claim 16, wherein the first ultrasound transducer array has a circular cross section having diameter between 1.5cm and 5cm.

20. A probe according to claim 16, wherein the inner ultrasound transducer array has an elliptical cross section whose maximum dimensions are: long axis between 2cm and 6 cm and short axis between 1.5cm and 5cm.

21. A probe according to claim 15, wherein fi is between 200KHz and 400KHz.

22. A probe according to claim 15, wherein f^ is double fi.

23. A probe according to claim 15, wherein some or all of the ultrasound beams have a sinusoidal waveform.

24. A probe according to claim 15, wherein fi is split into two close frequencies ft ± Αΐ, wherein Δί equals several KHz to several tens of KHz.

25. A probe according to claim 24, wherein Af equals between 2KHz and 25KHz.

26. A probe according to claim 15, wherein the second transducer array is comprised of several separate sub-elements that are operated at the same ultrasound frequency ft, but are driven separately with different electronic signal phases.

27. A probe according to claim 15, wherein the second transducer array is divided into two sections in order to enable spatial division of the ultrasound beam having frequency fi into two separate beams with different frequencies fi±Af.

28. A probe according to claim 26, wherein the separate sub-elements of which the second transducer array is comprised are divided into two sections in order to enable spatial division of the ultrasound beam having frequency fi into two separate beams with different frequencies fi±Af.

29. A probe according to claim 15, wherein the first ultrasound transducer array that operates at frequency Ϊ2 and which is already focused on the treated target volume is adapted to enable it to be operated in the Doppler CW or Pulse mode as a sensor probe for detection of the ultrasound echoes scattered by the bubbles.

30. A probe according to claim 15, comprising an additional ultrasound transducer designed to produce an ultrasound beam having a frequency higher than 2MHz and adapted to enable it to act as a Doppler detector element; wherein said additional transducer is designed to be focused on the same focal point as the other ultrasound transducers of said probe and is embedded within the probe in a position where it can be aligned in front of the interspaces between the ribs, in order to enable transmission and detection of the high frequency beam only through the inter-ribs space.

31. A system for non-invasive therapeutic treatment of subcostal lesions, said system comprising:

— a High Intensity Focused Ultrasound (HIFU) probe according to any one of claim 15 to claim 28; and

— a diagnostic ultrasound system, which is adapted to be used for realtime ultrasonic guidance, monitoring and control of the treatment.

32. A system for non-invasive therapeutic treatment of subcostal lesions, said system comprising:

— a High Intensity Focused Ultrasound (HIFU) probe according to either claim 29 or claim 30; and

33. A system for non-invasive therapeutic treatment of subcostal lesions, said system comprising:

34. A system according to claim 33 additionally comprising a position sensing system adapted to enable fusion of CT images of the lesion, which are taken with recognizable fiducial markers offline and prior to the HIFU treatment, with the 2D or 3D ultrasound images made with the 2D/3D diagnostic ultrasound system and with the focal zone monitoring, which are taken in real-time during the HIFU treatment.