Histotripsy using very short monopolar ultrasound pulses

Abstract

Apparatus and methods are provided for applying ultrasound pulses into tissue or a medium in which the peak negative pressure (P−) of one or more negative half cycle(s) of the ultrasound pulses exceed(s) an intrinsic threshold of the tissue or medium, to directly form a dense bubble cloud in the tissue or medium without shock-scattering. In one embodiment, a microtripsy method of Histotripsy therapy comprises delivering an ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse having at least a portion of a peak negative pressure half-cycle that exceeds an intrinsic threshold in the tissue to produce a bubble cloud of at least one bubble in the tissue, and generating a lesion in the tissue with the bubble cloud. The intrinsic threshold can vary depending on the type of tissue to be treated. In some embodiments, the intrinsic threshold in tissue can range from 15-30 MPa.

Description

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specification are herein incorporated by reference to the same extent as if each individual publication or patent application was specifically and individually indicated to be incorporated by reference.

FIELD

This disclosure generally relates to treating tissue with bubble clouds created by ultrasound therapy.

BACKGROUND

Histotripsy, or pulsed ultrasound cavitation therapy, is a technology where extremely short, intense bursts of acoustic energy induce controlled cavitation (microbubble formation) within the focal volume. The vigorous expansion and collapse of these microbubbles mechanically homogenizes cells and tissue structures within the focal volume. This is a very different end result than the coagulative necrosis characteristic of thermal ablation. To operate within a non-thermal, Histotripsy realm; it is necessary to deliver acoustic energy in the form of high amplitude acoustic pulses with low duty cycle.

Compared with conventional focused ultrasound technologies, Histotripsy has important advantages: 1) the destructive process at the focus is mechanical, not thermal; 2) bubble clouds appear bright on ultrasound imaging thereby confirming correct targeting and localization of treatment; 3) treated tissue appears darker (hypoechoic) on ultrasound imaging, so that the operator knows what has been treated; and 4) Histotripsy produces lesions in a controlled and precise manner. It is important to emphasize that unlike microwave, radiofrequency, or high-intensity focused ultrasound (HIFU), Histotripsy is not a thermal modality.

Histotripsy produces tissue fractionation through dense energetic bubble clouds generated by short, high-pressure, ultrasound pulses. Conventional Histotripsy treatments have used longer pulses from 3 to 10 cycles wherein the lesion-producing bubble cloud generation depends on the pressure-release scattering of very high peak positive shock fronts from previously initiated, sparsely distributed bubbles (the “shock-scattering” mechanism).

In conventional Histotripsy treatments, ultrasound pulses with ≥2 acoustic cycles are applied, and the bubble cloud formation relies on the pressure release scattering of the positive shock fronts (sometimes exceeding 100 MPa, P+) from initially initiated, sparsely distributed bubbles (or a single bubble). This has been called the “shock scattering mechanism”. This mechanism depends on one (or a few sparsely distributed) bubble(s) initiated with the initial negative half cycle(s) of the pulse at the focus of the transducer. A cloud of microbubbles then forms due to the pressure release backscattering of the high peak positive shock fronts from these sparsely initiated bubbles. These back-scattered high-amplitude rarefactional waves exceed the intrinsic threshold thus producing a localized dense bubble cloud. Each of the following acoustic cycles then induces further cavitation by the backscattering from the bubble cloud surface, which grows towards the transducer. As a result, an elongated dense bubble cloud growing along the acoustic axis opposite the ultrasound propagation direction is observed with the shock scattering mechanism. This shock scattering process makes the bubble cloud generation not only dependent on the peak negative pressure, but also the number of acoustic cycles and the amplitudes of the positive shocks. Without these intense shock fronts developed by nonlinear propagation, no dense bubble clouds are generated when the peak negative half-cycles are below the intrinsic threshold.

In more traditional ultrasound therapy, lesions are generated by forming an ultrasound beam into tight focal zones and using higher intensities within these foci to therapeutically modify tissue. Lesions are then generated either due to tissue heating (thermal therapy) or by mechanical agitation by energetic microbubble or bubble cloud formation (cavitation or Histotripsy). Minimum lesion size is typically determined by the diameter of the focal volume that, because of the limit due to diffraction, is around one wavelength of the propagating sound in the medium. A tightly focused ultrasound beam at 1 MHz (a typical therapeutic frequency) has a wavelength of about 1.5 mm in water or high water content tissue. Therefore, the focal diameter at 1 MHz is on the order of 1 mm or larger, depending on the transducer aperture. Using classical methods, getting smaller lesion diameters requires higher frequencies. Unfortunately, tissue penetration depth (due to increased tissue absorption and scattering) decreases with increasing frequency.

In the ultrasound cavitation based technology presented herein, much smaller diameters can be obtained at a given frequency where the lower limit in lesion size is the diameter of an individual microbubble, the primary effector in the process. Therefore, with proper pulse shaping and intensity, one can “beat” the “diffraction limit.”

SUMMARY OF THE DISCLOSURE

A method of treating tissue with ultrasound energy is provided, comprising the steps of delivering an ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse having at least a portion of a peak negative pressure half-cycle that exceeds an intrinsic threshold in the tissue to produce a bubble cloud of at least one bubble in the tissue, and generating a lesion in the tissue with the bubble cloud.

In some embodiments, the ultrasound pulse comprises at least one half cycle.

In one embodiment, the intrinsic threshold is greater than or equal to 15 MPa peak negative pressure. In another embodiment, the intrinsic threshold is approximately 28 MPa peak negative pressure. In one embodiment, the intrinsic threshold is between approximately 26 MPa and 30 MPa peak negative pressure.

In some embodiments, the ultrasound pulse has a frequency between approximately 0.1 MHz and 20 MHz.

In one embodiment, the portion of the peak negative pressure half-cycle that exceeds the intrinsic threshold produces the bubble cloud while a remaining portion of the peak negative pressure half-cycle that does not exceed the intrinsic threshold does not produce the bubble cloud.

In another embodiment, the ultrasound pulse comprises a first ultrasound pulse, the method further comprising delivering a second ultrasound pulse having a smaller amplitude than the first ultrasound pulse, such that a portion of the peak negative pressure half-cycle of the second ultrasound pulse that exceeds the intrinsic threshold is smaller than the portion of the peak negative pressure half-cycle of the first ultrasound pulse that exceeds the intrinsic threshold.

In one embodiment, the second ultrasound pulse produces a smaller bubble cloud than the first ultrasound pulse.

In some embodiments, the ultrasound pulse comprises a first ultrasound pulse, the method further comprising delivering a second ultrasound pulse having a larger amplitude than the first ultrasound pulse, such that a portion of the peak negative pressure half-cycle of the second ultrasound pulse that exceeds the intrinsic threshold is larger than the portion of the peak negative pressure half-cycle of the first ultrasound pulse that exceeds the intrinsic threshold.

In one embodiment, the second ultrasound pulse produces a larger bubble cloud than the first ultrasound pulse.

In some embodiments, the bubble cloud is formed in the tissue without shock-scattering.

A method of treating tissue with ultrasound energy is provided, comprising the steps of delivering an ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse having at least a portion of a peak negative pressure half-cycle that exceeds a peak negative pressure of 28 MPa in the tissue to produce a bubble cloud in the tissue, and generating a lesion in the tissue with the cavitation nuclei.

Another method of treating tissue with ultrasound energy is provided, comprising the steps of delivering a first ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse having at least a portion of a peak negative pressure half-cycle that exceeds an intrinsic threshold in the tissue to produce a first bubble cloud in the tissue, and delivering a second ultrasound pulse having a different amplitude than the first ultrasound pulse from the ultrasound therapy transducer into tissue to produce a second bubble cloud in the tissue, such that a portion of the peak negative pressure half-cycle of the second ultrasound pulse that exceeds the intrinsic threshold is different than the portion of the peak negative pressure half-cycle of the first ultrasound pulse that exceeds the intrinsic threshold.

An ultrasound therapy system is provided, comprising an ultrasound therapy transducer, and an ultrasound therapy generator coupled to the ultrasound therapy transducer, the ultrasound therapy generator configured to drive the ultrasound therapy transducer to deliver an ultrasound pulse into tissue, the ultrasound pulse having at least a portion of a peak negative pressure half-cycle that exceeds an intrinsic threshold in the tissue to produce a bubble cloud in the tissue.

In some embodiments, the ultrasound therapy generator further comprises a controller configured to generate complex waveforms to initiate the ultrasound pulse, a high voltage power supply coupled to the controller, an amplifier configured to receive and amplify the complex waveforms from the controller and high voltage power supply, and a matching network configured to match an impedance of the ultrasound therapy transducer to the amplifier.

In some embodiments, the intrinsic threshold is greater than or equal to 15 MPa peak negative pressure. In other embodiments, the intrinsic threshold is approximately 28 MPa peak negative pressure. In another embodiment, the intrinsic threshold is between approximately 26 MPa and 30 MPa.

In some embodiments, the ultrasound pulse has a frequency between approximately 0.1 MHz and 20 MHz.

In another embodiment, the portion of the peak negative pressure half-cycle that exceeds the intrinsic threshold produces cavitation nuclei while a remaining portion of the peak negative pressure half-cycle that does not exceed the intrinsic threshold does not produce cavitation nuclei.

BRIEF DESCRIPTION OF THE DRAWINGS

The novel features of the invention are set forth with particularity in the claims that follow. A better understanding of the features and advantages of the present invention will be obtained by reference to the following detailed description that sets forth illustrative embodiments, in which the principles of the invention are utilized, and the accompanying drawings of which:

FIG. 1 shows one embodiment of a Histotripsy therapy system.

FIG. 2 illustrates an intrinsic threshold in tissue.

FIGS. 3A-3B show a beam profile and an intrinsic threshold in tissue.

FIGS. 4A-4B show a beam profile with an adjusted amplitude and an intrinsic threshold in tissue.

FIGS. 5A-5C illustrate the production of a bubble cloud in tissue in response to adjusting an amplitude of subsequent ultrasound pulses.

DETAILED DESCRIPTION

Histotripsy is a noninvasive, cavitation-based therapy that uses very short, high-pressure ultrasound pulses to generate a dense, energetic, lesion-producing bubble cloud. This Histotripsy treatment can create controlled tissue erosion when it is targeted at a fluid-tissue interface and well-demarcated tissue fractionation when it is targeted within bulk tissue. Additionally, Histotripsy has been shown to be capable of fragmenting model kidney stones using surface erosion that is mechanistically distinct from conventional shockwave lithotripsy (SWL). Histotripsy therapy can be guided and monitored using ultrasound B-mode imaging in real-time, since 1) the bubble cloud appears as a temporally changing hyperechoic region in B-mode imaging, allowing the treatment to be precisely targeted, and 2) the echogenicity of the targeted region decreases as the degree of tissue fractionation increases, which can be used as a way of monitoring lesion production (image feedback) in real-time.

In this disclosure, methods, procedures, materials, and devices that produce ultrasound cavitation lesions much smaller than the diffraction limit (the minimal size of a focal zone . . . usually around one wavelength of the incident ultrasound) are presented. The technology is based on generating an extremely consistent bubble cloud with the first peak negative half-cycle of every single or multi-cycle ultrasound pulse. This can be accomplished with an ultrasound transducer capable of generating in the target volume a peak negative pressure exceeding an intrinsic threshold (sometimes called the “homogeneous” or de novo threshold) for cavitation, i.e., the level whereby the whole volume above this peak-negative threshold generates an immediate bubble cloud throughout the supra-threshold volume. The resulting bubble cloud can generate a therapeutic lesion in the tissue. Below this threshold, individual bubbles are sparsely generated, and a bubble cloud can only be developed with a complex interaction of succeeding cycles, usually by pressure reversal scattering of high peak positive half cycles off individual sparsely distributed seed bubbles generated by previous peak negative half cycles. This process produces a somewhat chaotic multilayer cloud with unpredictable dimensions of each cloud although usually very near the focal volume.

The de novo generated bubble cloud is much more consistent in spatial distribution. Moreover, the diameter of this de novo cloud is determined by the how much of the peak negative part of the waveform exceeds the intrinsic threshold. The amplitude of the waveform can be increased to cause more of the waveform to exceed the threshold, resulting in a larger bubble cloud diameter. The resulting tissue lesion diameter from an increase in the amplitude of the waveform also increases accordingly. Thus in some embodiments, this method, and appropriate hardware, allows a lesion diameter to be modified by changing the intensity of the ultrasound pulse. Moreover, if the pulse consists of only one cycle, with only one peak negative half cycle, pulses below the threshold will not generate bubble clouds because no succeeding positive half-cycles are available for the pressure-release scattering mode to generate a complete bubble cloud. In such circumstances (below de novo threshold), the first peak negative half-cycle would only generate sparsely distributed single bubbles.

Thus, the lower limit in size of a bubble cloud generated lesion is the diameter of a single bubble which is much smaller than the wavelength of the ultrasound waveform producing that bubble, therefore, the approach described herein, which can be referred to herein as “microtripsy”, is capable of generating lesion sizes much smaller than the diffraction limit determined by the frequency of the ultrasound source.

This disclosure provides mechanism for bubble cloud formation different than conventional shock-scattering bubble formation. Apparatus and methods are provided for applying ultrasound pulses into tissue or a medium in which the peak negative pressure (P−) of one or more negative half cycle(s) of the ultrasound pulses exceed(s) an intrinsic threshold of the tissue or medium, to directly form a dense bubble cloud in the tissue or medium without shock-scattering. This intrinsic threshold can be defined by a very sharp transition zone, and can be relatively insensitive to the inhomogeneities in soft tissue or the lack of dissolved gases.

Moreover, this intrinsic threshold may be independent of the applied positive pressure, since the influence of shock-scattering is minimized when applying pulses less than 2 cycles. In some instances, the intrinsic threshold, where there is a high probability of cavitation (formation of a dense bubble cloud) for one single pulse is between 26.4-30.0 MPa in samples with high water content, including water, hydrogel, and soft tissue. In tissue with a higher fat content, including fatty tissues and brain tissue, the intrinsic threshold can be as low as 15 MPa. In one specific embodiment, illustrated by FIG. 2, the intrinsic threshold for a high probability of cavitation in tissue can be approximately 28 MPa. In the chart of FIG. 2, it can be seen that bubbles in this size range display a sharp threshold around a peak negative pressure of ˜28 MPa. Because of the small size of the microbubbles, the cavitation threshold is nearly independent of the frequency range of interest for ultrasound therapy. For example, an intrinsic threshold with a high probability of cavitation has been identified clinically at 28.1 MPa at a frequency of 0.1 MHz, compared to a threshold of 28.4 MPa at 10 MHz.

Because of the sharp, highly repeatable threshold, at negative pressure amplitudes not significantly greater than this, a dense energetic lesion-forming cloud of microbubbles can be generated consistently with a spatial pattern similar to the part of the negative half cycles(s) exceeding the intrinsic threshold. The part of the therapy pulse exceeding intrinsic threshold, and the resulting spatial extent of the lesion forming bubble cloud, are independent of positive shocks and the somewhat chaotic shock-scattering phenomenon. Therefore, the spatial extent of the lesion generated using this mechanism is expected to be well-defined and more predictable. The lesion size can be easily controlled, even for small lesions, simply by a precise adjustment of the therapy pulse amplitude.

Apparatus and methods described herein allow high tissue penetration with minimal aberration and signal degradation while generating acceptably precise lesions. Since clinical targets shadowed by bone are highly attenuating and aberrating, some applicable targets are those reached in the brain (shadowed by skull) or thoracic cavity (shadowed by ribs), although transmission through other complex aberrating tissues is possible, e.g., for trans-uterine procedures.

Lower frequencies propagate much more cleanly through complex media, minimizing aberrations and attenuation, but produce larger lesions, based on the diffraction limit dictating larger focal zones, when conventional ultrasonic therapeutic modalities are used. By using the methods and apparatus described herein, lower frequencies can be used while, at the same time, maintaining small and precise lesion sizes.

FIG. 1 illustrates a Histotripsy system configured to generate cavitation bubbles or bubble clouds in tissue according to the methods and embodiments described herein. A Histotripsy system and generator is configured to generate complex waveforms in order to support the ultrasound pulse sequences described herein. A simplified block diagram of system 100 is shown in FIG. 1. The main components of the system are: Computer/controller 102, USB to Serial Converter 104, Microcontroller 106, FPGA (Field Programmable Gate Array) 108, High Voltage Controller and Power Supply 110, Amplifier 112, and Therapy Transducer 114.

All controls for the generator can be established using a “Histotripsy Service Tool” software that can run on the computer/controller 102 (e.g., a standard PC, laptop, tablet, or other electronic computing system) and communicates to the generator via a connector such as USB serial communication 104. Therapy, including bubble cloud cavitation and tissue ablation can be imaged and tracked in real time using an imaging system 115, such as an ultrasound imaging system or probe.

The system 100 can be configured to receive multiple sets of different driving parameters and loop them, which give the ability to the user to create wide range of custom sequences where all parameters (pulse repetition frequency (PRF), voltage amplitude, number of cycles, number of pulses per set, frequency, transducer element channels enabled, and time delays) can be set differently for every pulse generated. Time delays between pulses can be specified by the PRF for a parameter set or by specifying zero as the number of cycles per pulse.

For overall voltage amplitude regulation, level of high voltage can be changed accordingly through the Microcontroller 106 and HV Controller 110. This method cannot be used for dynamic voltage amplitude changes between two pulses since it will take too long for all capacitors on the HV line to discharge. For dynamic voltage amplitude changes between pulses, PWM (pulse width modulation) can be used at the FPGA 108 where the duty cycle of the pulse may be modulated in order to produce the desired pulse voltage and resultant pressure amplitude.

Histotripsy Service Tool

Histotripsy Service Tool is an application that can be run on any PC or computing system and may be used for controlling the system. The Histotripsy Service Tool can start/stop the therapy, set and read the level of high voltage, therapy parameters (PRF, number of cycles, duty ratio, channel enabled and delay, etc), and set and read other service and maintenance related items.

USB to Serial Converter

USB to Serial converter 104 can convert USB combination to serial in order to communicate to the Microcontroller 106. It should be understood that other converters (or none at all) may be used in embodiments where the connection between the generator and the controller is not a USB connection.

Microcontroller

The Microcontroller 106 communicates to the computer/controller 102 (Histotripsy Service Tool) to set/read working parameters, start/stop the therapy, etc. It can use internal flash memory or other electronic storage media to store all the parameters. The Microcontroller communicates to the FPGA 108 all driving parameters that are necessary to generate complex pulsing. It also communicates using serial communication or other electronic communication to the high voltage controller and power supply 110 where it can set/read the proper level of driving voltage.

FPGA

The FPGA 108 receives the information from the Microcontroller 106 and it can generate the complex pulsing sequence that is required to drive the amplifier 112. The FPGA can run on 100 MHz clock since speed of pulsing is critical to be timed in at least 10 ns increments.

High Voltage Controller and Power Supply

The High Voltage Controller and Power Supply 110 receives the commands from the Microcontroller 106 regarding the level of DC voltage that needs to be supplied to the amplifier circuitry in order to have an adequate voltage amplitude level at the output of the amplifier.

Amplifier

The Amplifier 112 receives pulses generated by the FPGA and is supplied with high voltage from High Voltage Controller and Power Supply. It generates high voltage amplitude pulses that are fed to the Therapy Transducer 114 through the matching network components which properly matches the impedance of the therapy transducer to the impedance of the amplifier. It can be necessary to use a large number of capacitors that can store enough energy to support peak current demand during the generation of high voltage amplitude pulses.

Therapy Transducer

The Therapy Transducer 114 can be a single or multi-element ultrasound therapy transducer configured to generate and deliver the ultrasound therapy pulses described herein into tissue or other mediums. The active transducer elements of the Therapy Transducer can be piezoelectric transducer elements. In some embodiments, the transducer elements can be mounted to an acoustic lens with a common geometric focus.

Imaging System

The system can include an imaging system, such as an ultrasound imaging system, to monitor therapy and track cavitation and tissue ablation in real time. The cavitating bubble cloud generated according to the methods and system of this disclosure can appear as a temporally changing hyperechoic region in B-mode imaging, allowing the treatment to be precisely targeted.

The probability for the generation of a dense bubble cloud using a single, short pulse (≤2 cycles) is a function of the applied peak negative pressure P−. This cavitation probability curve follows a sigmoid function, given by

$\begin{array}{cc}{P}_{\mathrm{cav}}\left(p_{-}\right)=\frac{1}{2}\left[1+\mathrm{erf}\left(\frac{p_{-}-pt}{\sqrt{2{\sigma }^2}}\right)\right],& \left(1\right)\end{array}$

where erf is the error function, p_t is the pressure that gives a cavitation probability (P_cav) of 0.5, and σ is a variable that relates to the transition width in the cavitation probability curve, with ±σ giving the difference in pressure for cavitation probability from 0.15 to 0.85. The p_t's and σ's in water, gelatin gel, and high water content soft tissue were found to be within the range of 26-30 MPa (p_t) and 0.8-1.4 MPa (σ).

An intrinsic pressure threshold is surprisingly similar for cavitation in water, tissue-mimicking phantoms, and tissue. The occurrence of cavitation is a stochastic process; it may or may not occur for a single pulse with finite probability, even though all parameters of the system are identical pulse to pulse. At low pressures, the probability is almost zero and with large enough pressure amplitude the probability is 1 (i.e. cavitation always happens somewhere in the focus). A key factor for microtripsy is that the intrinsic threshold is fairly sharp, meaning there is a small pressure difference between cavitation probabilities of 0 and 1. For instance, in blood or tissue, the probability is near 0 at 25 MPa peak negative pressure and nearly 1 at ˜28 MPa peak negative pressure.

When a pulse passes through an ultrasound transducer focus, different positions in the medium experience different pressures. The center of the focus experiences the greatest pressure amplitude and the focal volume and dimensions are usually determined by the region which experiences peak negative pressures greater than some value relative to the maximum. In our case, we can create a map of the probability of cavitation occurring vs. position for a given pulse. Since there is a small difference in pressure between where cavitation occurs with high probability to where it occurs with very low probability, the boundary is very sharp. If the pressure output is chosen appropriately so that only the center of the focus is just above the threshold, then cavitation is likely to occur only at the focus center, a region which can be considerably smaller than 1 acoustic wavelength. In turn, it is possible that mechanical ablation of tissue smaller than 1 acoustic wavelength can be performed.

To achieve this sort of precision, short-duration (as low as ½ cycle) high peak negative pressure pulses can be generated and focused with the system described above. As described, the peak negative pressure which can be produced at the focus must exceed the intrinsic threshold to produce cavitation. Using long pulses can cause a complex interaction between a few existing microbubbles and the ultrasound pulse. This pulse will scatter from the bubbles and amplify the local pressure, causing a greater number of cavitation bubbles to expand. This limits the possible precision of microtripsy, and is necessarily circumvented by only using short pulses. Transducers to generate this sort of pressure pulse must be highly focused with high bandwidth. This output also requires an electronic driver or amplifier capable of providing the necessary high voltage to the transducer in short pulses.

As described above, an intrinsic threshold in tissue for bubble cavitation can be approximately 28 MPa. For cavitation nuclei in a size range on the order of 2.5 nm, surface tension is the dominant force controlling the threshold by the Laplace pressure. In this respect, it is similar to the Blake threshold. Systems and methods herein can include a frequency range of interest for therapy between 0.1-20 MHz. While the threshold of an individual nucleus may not be a strong function of frequency, the probability of cavitation also depends on the number of nuclei within the focal volume. As such, if an equivalent transducer with a higher frequency were applied, one would expect an incremental increase in the predicted pressure threshold due to the smaller volume of the focus. It appears that a greater negative pressure excursion will activate additional smaller nuclei coexisting with those which cavitate near the 28 MPa threshold.

Systems and methods of producing cavitation in tissue utilizing a “microtripsy” technique will now be described. An ultrasound therapy transducer, such as the transducer and Histotripsy system described in FIG. 1, can be configured to deliver one or more ultrasound pulses into tissue towards a focal point of the transducer. The one or more ultrasound pulses can include at least one negative half cycle. The system is configured to generate cavitation bubble(s) in tissue when the peak negative pressure of the waveform exceeds the intrinsic threshold in the tissue.

In some embodiments, the waveform can be generated such that only a portion of the negative half cycle has a negative pressure exceeding the intrinsic threshold. The portion of the waveform that includes a peak negative pressure greater than the intrinsic threshold in tissue will generate a bubble cloud in the tissue. Thus, the amplitude of the waveform can be adjusted to control the size of the bubble-cloud generated in the tissue. For example, the amplitude of the waveform can be reduced so that only a small portion, or the peak of the waveform exceeds the intrinsic threshold, so as to reduce the size of lesions formed in tissue from the bubble cloud. Alternatively, the amplitude of the waveform can be increased so that all or nearly all of the waveform includes a peak negative pressure that exceeds the intrinsic threshold, thereby creating a larger bubble cloud thereby increasing the size of the lesion in the tissue.

In some embodiments, the ultrasound pulse(s) can have a frequency ranging from 0.1-20 MHz or more. Reducing the frequency, such as applying ultrasound pulses to tissue at 0.1 MHz, facilitates the formation of smaller lesions in tissue with the resulting bubble cloud. Reduced frequencies an also allow bubble cloud formation deeper into tissue due to reduced attenuation.

The ultrasound pulse(s) delivered to tissue can have as few as ½ cycles, as long as at least one of the peak negative pulses exceeds the intrinsic threshold of the tissue. In one embodiment, the Histotripsy system and transducer can be configured to produce and deliver only a monopolar peak negative pulse into the tissue wherein the peak negative pulse exceeds the intrinsic threshold of the tissue.

FIGS. 3A-3B illustrate a beam profile of an ultrasound pulse 316 having a negative half cycle. In the illustrated example, the pulse has a negative pressure ranging from approximately 5 MPa to a peak of approximately 50 MPa. FIG. 3A also shows the intrinsic threshold 318 at which cavitation can be formed in tissue. In this embodiment, the intrinsic threshold is approximately 28 MPa. Referring to FIG. 3A, it can be understood that the portion of the pulse that includes a peak negative pressure exceeding the intrinsic threshold 318 is configured to generate cavitation bubble(s) in the tissue with a high probability. This portion of the pulse that exceeds the intrinsic threshold is illustrated between points 320 and 322.

FIG. 3B shows how the resulting lesion size from cavitation generated by the beam profile can be estimated. In FIG. 3B, a width 324 of the lesion can be estimated by taking the distance between the two points of the beam profile that exceed the intrinsic threshold. Thus in this example, the pulse with a peak negative pressure that exceeds the intrinsic threshold between points 320 and 322 can result in a lesion size estimated to be approximately 2 mm in width.

FIGS. 4A-4B illustrate how adjusting the amplitude of the ultrasound pulse can adjust the amount of cavitation produced, by changing the amount of the waveform that has a peak negative pressure exceeding the intrinsic threshold in tissue. In FIG. 4A, the ultrasound pulse 416 is shown having a negative pressure ranging from approximately 10 MPa to 80 MPa with the intrinsic threshold 418 shown at approximately 28 MPa. In this example, a large portion of the pulse 416 has a negative pressure that exceeds the intrinsic threshold. In FIG. 4B, the amplitude of a subsequent waveform has been decreased to produce a pulse having a negative pressure ranging from approximately 4 MPa to 30 MPa. In this example, only a small portion, or the very tip of the pulse, has a negative pressure that exceeds the intrinsic threshold. Reducing the amplitude of the pulse reduces the amount of the waveform that generates a bubble cloud in tissue, allowing for a smaller and more controlled bubble cloud and thus, a smaller and more controlled lesion in tissue. It should be understood that the amplitudes of subsequent waveforms can also be increased to increase the portion of the waveform that generates a bubble cloud in tissue, allowing for a larger bubble cloud and a larger lesion in tissue.

FIGS. 5A-5C illustrate how adjusting the amplitude or adjusting the portion of the negative half-cycle that exceeds the intrinsic threshold in tissue can affect the amount and size of the bubble cloud and resulting lesion in tissue. As described above, the shape of the bubble cloud generated by the intrinsic threshold mechanism is well-confined and corresponds well to the shape of the transducer focal zone. For illustrative purposes, the bubble clouds 526 shown in FIGS. 5A-5C correspond to a transducer focal zone in the shape of an ellipse or a “football” shape. FIG. 5A can illustrate a baseline bubble cloud 526. In FIG. 5B, the amount and size of the bubble cloud 526 can be increased compared to FIG. 5A by increasing an amplitude of the ultrasound pulse to cause a larger portion of the negative half-cycle to exceed the intrinsic threshold. In contrast, in FIG. 5C, the amount and size of the bubble cloud 526 can be decreased compared to FIG. 5A by decreasing an amplitude of the ultrasound pulse to cause a smaller portion of the negative half-cycle to exceed the intrinsic threshold. The resulting bubble cloud in FIG. 5C shows only a few bubbles generated in tissue.

In some embodiments, microtripsy therapy including the generation of cavitation and ablation of tissues can be monitored under real-time imaging, such as under ultrasound imaging. The real-time imaging can be used to monitor the bubble cloud formation, and also used as feedback to adjust the size of the bubble cloud. For example, in one embodiment, an initial ultrasound pulse having a peak negative pressure that exceeds the intrinsic threshold can be transmitted into tissue to generate a bubble cloud. The size of the bubble cloud can be monitored in real time with the imaging system. Next, a second ultrasound pulse with an amplitude different than the initial ultrasound pulse can be transmitted into the tissue to generate a bubble cloud. If the amplitude of the second ultrasound pulse is less than the initial ultrasound pulse but the pulse still includes a negative pressure above the intrinsic threshold, then the bubble cloud generated will be smaller than was generated with the initial pulse. If the amplitude of the second pulse is larger, then the bubble cloud generated will be greater than with the initial pulse.

Precise and controlled lesions can be generated in tissue using the intrinsic threshold mechanism of Histotripsy therapy described herein. In comparison to the bubble clouds and lesions generated by a shock-scattering mechanism, those generated by the intrinsic threshold mechanism have two advantageous characteristics. First, the shape of the bubble cloud generated by the intrinsic threshold mechanism is well-confined and corresponds well to the shape of the transducer focal zone, whereas the shape of the bubble cloud generated by the shock scattering mechanism is variable and somewhat unpredictable. Because tissue fractionation is directly correlated to the activity of the bubble cloud, the shape of the lesions produced by the intrinsic threshold mechanism is more predictable and has a better agreement to the shape of the transducer focal zone. For waveforms with a peak negative pressure near the intrinsic threshold pressure, single bubbles appear primarily in a region near the center of the focal zone. When the focal pressure is significantly in excess of the cavitation threshold, cavitation bubbles tend to occupy a region similar in shape to the focal zone for each pulse.

Second, the bubble clouds generated by the intrinsic threshold mechanism are more uniform and consistent within the region exceeding the intrinsic threshold, whereas the bubble clouds generated by the shock-scattering mechanism can be isolated to subvolumes of the focus due to complex scattering behavior. Therefore, the intrinsic threshold mechanism can potentially lead to more efficient and complete lesion development in tissue.

Moreover, using this intrinsic threshold mechanism, very small and controlled lesions can be generated by allowing only a small fraction of the focal region to exceed the cavitation threshold. For example, experiments using this method have resulted in lesions with diameters of less than 1 mm. These values were much smaller than the wavelength of the transducer (λ=˜3 mm for 500 kHz) used to generate the lesions. These results demonstrate that microscopic and precise lesions can be achieved using the intrinsic threshold mechanism, which can also be referred to as “microtripsy” herein.

This “microtripsy” technique based on the bubble clouds generated by the intrinsic threshold mechanism can be quite beneficial in the case where microscopic and well-defined tissue ablation is required. Especially, very precise lesions can still be achieved using a low frequency transducer by allowing only a small fraction of the focal zone to exceed the intrinsic threshold. Low frequency transducers would be favorable in applications that require long ultrasound penetration depth or where the intended targets have very attenuative overlying tissues, such as in transcranial brain therapy. Moreover, low-frequency single cycle (or close) pulses minimize phase aberration

This “microtripsy” procedure uses the highly repeatable and very sharp transition zone in cavitation probability inherent to bubble cloud generation above the intrinsic threshold. Lesion sizes in both axial and lateral directions can be increased by increasing the applied peak negative pressure. These lesion sizes corresponded well to the dimensions of the focal beam profile estimated to be beyond the intrinsic cavitation threshold. This “microtripsy” technique can be significantly useful in the clinical applications where precise, microscopic tissue ablation is required, particularly where low frequencies are indicated while still maintaining small precise sub-wavelength lesions.

The aim of a “microtripsy” system can be to generate very short (e.g., as low as one half cycle) high intensity ultrasound pulses wherein the peak negative pressure of the pulse exceeds the intrinsic (or “homogeneous”) threshold in the target media, such as tissue. Therefore, systems, apparatus, and methods described herein are configured to generate bubble clouds of precise and repeatable spatial distribution the diameter, which can be smaller than a wavelength of the ultrasound propagating wave inducing the bubble cloud, i.e., capable of producing lesion diameters “below” the diffraction limit.

The technology described herein allows for applications beyond the physical limitations of current thermal or cavitation based therapeutic modalities. In particular, many advantages derive from employing much lower frequencies while keeping the diameter of generated bubble clouds low (below the diffraction limit). These include: 1. Higher penetration depth, 2. Lower absorption, 3. Lower thermal complications due to heating, 4. Reduced aberration levels, especially through complex tissues including bone, and, 5. Reduced need for aberration correction methods. This can all be accomplished without suffering the usual downside (at lower frequencies) of a larger focal volume and a larger lesion.

The combination of these, and other advantages to be discussed below, will allow targets to be addresses that are currently extremely difficult by noninvasive ultrasonic means. These targets include, but are not limited to: transcranial (through the skull) noninvasive therapeutic procedures including precise lesion formation, immediate (rapid) and noninvasive treatment of stroke inducing blood clots in the brain, brain tumors, and other neural disorders treated by very precise disruption or lesion formation in neural tissues, and precise disruption of the blood brain barrier for drug delivery across drug diffusion-resistant protective membrane systems; trans-thoracic non-invasive procedures through and around the ribs. The largest application can include the treatment of liver mets, direct cardiac ablation for treatment of arrhythmias, treatment of tumors in the spleen, etc. Other applications are considered such as pediatric applications (HLHS, Trans-uterine, etc.) or thrombolysis in small vessels.

All these targets can be reached by using lower frequency ultrasound without giving up precise localization and small lesion size. Moreover, most of the treatments can be delivered by single focus inexpensive transducers without the need for the complex and expensive phased array transducers necessary for aberration correction processes. And, of course, therapy can now be delivered without the expensive means for quantifying the spatial distribution of the acoustic aberrations between the transducer and the target volume (MRI machines, invasive acoustic hydrophones or beacons and associated electronics to determine the quantitative spatial distribution of the acoustic aberrations). The methods disclosed herein are remarkably aberration resistant even not accounting for the possibility, with microtripsy, of using much lower frequencies.

With knowledge of the bubble-cloud dynamics, the pressure thresholds, and the beam pressure profiles for the transducer used for therapy, it can be possible to create a more predictive mechanical ultrasound ablative therapy with cavitation than has been previously achieved. The typical shock-scattering mechanism for forming bubble clouds, which uses multi-cycle pulses, has some limitation, in that it depends on the locations of inhomogeneties and stabilized gas bubbles in the medium to act as single bubble scatterers prior to growing the dense cloud which causes the majority of tissue disruption. Unfortunately, it is very difficult to determine the location of heterogeneous nuclei before treatment, and this can lead to clouds forming in distinct locations in the focal zone while other locations remain completely uncavitated. In contrast, the single pressure excursion pulse described herein can generate cavitation uniformly within the focal zone. While cavitation has generally been regarded as unpredictable in ultrasound therapy, the pulsing regime of this disclosure can create a very controllable situation for planning therapy. Around the cavitation threshold, single bubbles can be generated by a single pulse. In contrast to thermal ultrasound therapy, where applying too great of an acoustic dose can cause damage to collateral tissue by diffusion of heat, the damage with histotripsy can be limited to the focal region and overtreating may have few negative consequences, as the spatial boundary between very high probability and low probability is very narrow because of the intrinsically small value for the probability transition zone σ.

The advantages of bubble cloud generation with as little as a half cycle (or just a few cycles) high intensity pulses with peak negative pressures exceeding the de novo (or homogeneous) threshold are as follows:

The bubble cloud extent can be controlled by the fraction of the peak negative waveform exceeding the intrinsic threshold, which is proportional to pulse intensity. Therefore, the minimum lesion size is limited only by the size of the smallest bubble cloud possible which is the diameter of a single bubble. Bubbles formed near the threshold for a 1 MHz transducer are 100-200 m diameter in tissue, which is 0.06-0.13 wavelengths. Bubbles will grow inversely proportionate to frequency, but their fraction of a wavelength will be similar for most cases. Tissue confinement can limit the maximum expansion of bubbles below these values as well. Therefore, instead of the diffraction limit of several millimeters, the minimum lesion diameter may be as small as 100 micrometers, or a reduction in lesion diameter of an order of magnitude below the diffraction limit.

Smaller bubble clouds mean, for a given fixed lesion diameter, much lower frequencies can be used increasing depth of penetration and decreasing absorption and scattering.

Lower absorption means greatly reduced heating and unwanted hot spots, particularly at tissue boundaries. Such undesired (and unpredictable) thermal consequences often define the limits for certain ultrasound therapies at depth.

Tissue aberrations decrease with frequency allowing much cleaner beams to be formed through complex tissues reducing the need for aberration correction.

The “microtripsy” approach uses the portion of the ultrasound waveform above the intrinsic threshold to generate bubble clouds, thus allowing a large degree of protection against adverse effects of high ultrasound beam side-lobes.

Lower aberrations levels (from the allowed lower frequencies) and the relative insensitivity to aberrations will allow greatly reduced overall system expense by greatly reducing the need for expensive aberration correction hardware and software.

Lower frequencies and lack of need for aberration correction schemes will allow clinical targets like the brain and liver to be treated through the skull and ribs, respectively, without an increase in lesion size.

The de novo generated bubble clouds are remarkably consistent in size for single cycle pulses, and the size can be adjusted by adjusting the pulse amplitude without a need to change the source frequency or the source aperture. As stated several times before, this bubble cloud diameter, and resulting lesion diameter, can be much below the diffraction limit in size.

When bubble clouds are generated by de novo single pulse Histotripsy, the treated volume is remarkably consistent from pulse to pulse with the bubbles forming at random (not predetermined) locations in the focal volume. In other approaches to generating bubble clouds, the location of individual bubbles can depend on memory effects from previous pulses (due to persistent bubble remnants or micronuclei). This memory effect often results in clouds being generated non-homogeneously in the focal volume with some sub-volumes generating no bubbles at all. This results in non-homogeneous lesions in the focal volume with islands of untreated tissues. Many additional pulses are necessary to expand these treated zones to include the un-homogenized islands. Although there are active and passive ways to avoid this memory effect, they require greater complexity.

The zone raised above the de novo threshold (the “de novo volume”) generates bubbles throughout this volume regardless of pre-existing nuclei due to the memory effect, or due to natural dissolved gas, or other cavitation nuclei, in the target volume. Because the de novo threshold is higher than for other forms of cavitation, bubble clouds in the de novo volume tend to be remarkably consistent from pulse to pulse and in different environments.

Because the optimum “microtripsy” pulse comprises at least a single half cycle, the thermal potential of this therapy is minimized due to the very short pulses.

Because the de novo volume is resistant to memory and other pre-existing bubble clouds, much greater pulse repetition frequencies (PRFs) are possible before remnants of previous bubble clouds interfere in a destructive manner in the generation of the latest bubble cloud, a phenomenon common at higher PRFs in other Histotripsy approaches.

Short single cycle pulses are more resistant to aberrations because less time is available for multi-path scattered signals to form up constructively to produce higher side-lobes.

Since the bubble clouds in the de novo volume are not dependent on pressure release scattering from previous bubbles randomly generated by separate multi-cycle bursts or succeeding cycles in a given burst, all below the de novo threshold, the need for more than one initiation pulse (or burst) to get a fully formed bubble cloud disappears.

The data structures and code described in this detailed description are typically stored on a computer-readable storage medium, which may be any device or medium that can store code and/or data for use by a computer system. The computer-readable storage medium includes, but is not limited to, volatile memory, non-volatile memory, magnetic and optical storage devices such as disk drives, magnetic tape, CDs (compact discs), DVDs (digital versatile discs or digital video discs), or other media capable of storing computer-readable media now known or later developed.

The methods and processes described in the detailed description section can be embodied as code and/or data, which can be stored in a computer-readable storage medium as described above. When a computer system reads and executes the code and/or data stored on the computer-readable storage medium, the computer system performs the methods and processes embodied as data structures and code and stored within the computer-readable storage medium.

Furthermore, the methods and processes described above can be included in hardware modules. For example, the hardware modules can include, but are not limited to, application-specific integrated circuit (ASIC) chips, field-programmable gate arrays (FPGAs), and other programmable-logic devices now known or later developed. When the hardware modules are activated, the hardware modules perform the methods and processes included within the hardware modules.

The examples and illustrations included herein show, by way of illustration and not of limitation, specific embodiments in which the subject matter may be practiced. As mentioned, other embodiments may be utilized and derived there from, such that structural and logical substitutions and changes may be made without departing from the scope of this disclosure. Such embodiments of the inventive subject matter may be referred to herein individually or collectively by the term “invention” merely for convenience and without intending to voluntarily limit the scope of this application to any single invention or inventive concept, if more than one is, in fact, disclosed. Thus, although specific embodiments have been illustrated and described herein, any arrangement calculated to achieve the same purpose may be substituted for the specific embodiments shown. This disclosure is intended to cover any and all adaptations or variations of various embodiments. Combinations of the above embodiments, and other embodiments not specifically described herein, will be apparent to those of skill in the art upon reviewing the above description.

Claims

1. A method of treating tissue with ultrasound energy, comprising the steps of: delivering an ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse comprising only a monopolar peak negative pulse having a peak negative pressure that exceeds an intrinsic threshold in the tissue to produce a bubble cloud of at least one bubble in the tissue; andgenerating a lesion in the tissue with the bubble cloud.

2. The method of claim 1, wherein the ultrasound pulse comprises one half cycle.

3. The method of claim 1 wherein the intrinsic threshold is greater than or equal to 15 MPa peak negative pressure.

4. The method of claim 1 wherein the intrinsic threshold is 28 MPa peak negative pressure.

5. The method of claim 1 wherein the intrinsic threshold is between 26 MPa and 30 MPa peak negative pressure.

6. The method of claim 1 wherein the ultrasound pulse has a frequency between 0.1 MHz and 20 MHz.

7. The method of claim 1 wherein the ultrasound pulse comprises a first ultrasound pulse, the method further comprising delivering a second ultrasound pulse having a smaller amplitude than the first ultrasound pulse, such that a peak negative pressure of the second ultrasound pulse exceeds the intrinsic threshold but is smaller than the peak negative pressure of the first ultrasound pulse.

8. The method of claim 7, wherein the second ultrasound pulse produces a smaller bubble cloud than the first ultrasound pulse.

9. The method of claim 1 wherein the ultrasound pulse comprises a first ultrasound pulse, the method further comprising delivering a second ultrasound pulse having a larger amplitude than the first ultrasound pulse, such that a peak negative pressure of the second ultrasound pulse exceeds the intrinsic threshold but is larger than the peak negative pressure of the first ultrasound pulse.

10. The method of claim 9, wherein the second ultrasound pulse produces a larger bubble cloud than the first ultrasound pulse.

11. The method of claim 1, wherein the bubble cloud is formed in the tissue without shock-scattering.

12. A method of treating tissue with ultrasound energy, comprising the steps of: delivering an ultrasound pulse from an ultrasound therapy transducer into tissue, the ultrasound pulse comprising only a monopolar peak negative pulse having a peak negative pressure that exceeds 28 MPa in the tissue to produce a bubble cloud in the tissue; andgenerating a lesion in the tissue with the bubble cloud.