The present invention relates, generally, to systems and methods for ultrasound focusing. In particular, various embodiments are directed to efficient methods of focusing a phased array of ultrasound transducer elements, using preparatory measurements to adjust the relative phases of the transducer elements.
Thermal ablation, as may be accomplished using focused ultrasound, has particular appeal for treating tissue within the brain and other tissue regions deep within the body, because it generally does not disturb intervening or surrounding healthy tissue. Focused ultrasound may also be attractive, because acoustic energy generally penetrates well through soft tissues, and ultrasonic energy, in particular, may be focused towards focal zones having a cross-section of only a few millimeters due to relatively short wavelengths (e.g., as small as 1.5 millimeters (mm) in cross-section at one Megahertz (1 MHz)). Thus, ultrasonic energy may be focused at a region deep within the body, such as a cancerous tumor or other diseased tissue, to ablate the diseased tissue without significantly damaging surrounding healthy tissue.
To focus ultrasonic energy towards a desired target, a piezoelectric transducer may be used that includes a plurality of transducer elements. A controller may provide drive signals to each of the transducer elements, thereby causing the transducer elements to transmit acoustic energy such that constructive interference occurs at a “focal zone”. The focal zone is typically defined as the region of intensity higher than half maximum, and is commonly characterized by a “peak width” in a given direction. The peak width may be anisotropic. In fact, most realized instrumental systems produce an elliptical shaped peak cross-section at half maximum. At the focal zone, sufficient acoustic energy may be delivered to generate the desired tissue activation (e.g., heating, necrosis, neural stimulation, etc . . . ) within the focal zone and for a sufficient period until tissue affects occurs. Preferably, tissue along the path through which the acoustic energy passes (“the pass zone”) outside the focal zone, is affected (e.g, heated) only minimally, if at all, thereby minimizing damaging tissue outside the focal zone.
Phased arrays of ultrasound transducers are well-known as a system for focusing ultrasound energy at target sites inside the body. Constructive and destructive interference of acoustic waves transmitted by multiple transducers can be used to deliver complex spatiotemporal patterns of acoustic waves. Generally, phased arrays use tens to hundreds or even thousands of ultrasound transducers distributed spatially on the surface of the body. For instance, a phased array placed on the head can be used to target an area deep in the brain. However, phased arrays have important limitations for delivering ultrasound transcranially for neuromodulation. Phased arrays use spatially distributed transducers, requiring a larger form factor. Moreover, large and generally unportable power and control components are required to manage the timing, intensity, phase, and other properties of the ultrasound waves transmitted by each of the transducers.
The prior art of ultrasound focusing onto tissue in general and brain tissue in particular is exemplified in USA patents U.S. Pat. No. 8,932,237, U.S. Pat. No. 5,329,930, U.S. Pat. No. 4,817,614, U.S. Pat. No. 8,088,067, U.S. Pat. No. 6,128,958, U.S. Pat. No. 7,611,462, U.S. Pat. No. 5,984,881, and references therein, the entire disclosures of which are hereby incorporated by reference.
Focused ultrasound (i.e., acoustic waves having a frequency greater than about 20 kilohertz) can be used to image or therapeutically treat internal body tissues within a patient. For example, ultrasonic waves may be used to ablate tumors, eliminating the need for the patient to undergo invasive surgery. For this purpose, a piezo-ceramic transducer is placed externally to the patient, but in close proximity to the tissue to be ablated (the “target”). The transducer converts an electronic drive signal into mechanical vibrations, resulting in the emission of acoustic waves (a process hereinafter referred to as “sonication”). The transducer may be shaped so that the waves converge in a focal zone. Alternatively or additionally, the transducer may be formed of a plurality of individually driven transducer elements whose phases (and, optionally, amplitudes) can each be controlled independently from one another and, thus, can be set so as to result in constructive interference of the individual acoustic waves in the focal zone. Such a “phased-array” transducer facilitates steering the focal zone to different locations by adjusting the relative phases between the transducers, and generally provides the higher a focus quality and resolution, the greater the number of transducer elements. Magnetic resonance imaging (MRI) may be utilized to visualize the focus and target in order to guide the ultrasound beam.
While the transducer is located external to the patient, it must be in direct contact and tightly coupled with a media that efficiently transmits the high frequency ultrasound waves. For example, the transducer can be positioned in a liquid bath that is capable of efficient transmission of the ultrasound waves. The patient's body must also be wetted and tightly coupled to the transmission media in order to ensure an optimal acoustic wave transmission path from the transducer to the focal zone.
While system parameters are generally fixed for a given transducer array, tissue homogeneity may vary significantly from patient to patient, and even between different tissue regions within the same patient. Tissue inhomogeneity may decrease intensity of the acoustic energy at the focal zone and may even move the location of the focal zone within the patient's body. Specifically, because the speed of sound differs in different types of tissue, as portions of a beam of acoustic energy travel along different paths towards the focal zone, they may experience a relative phase shift or time delay, which may change the intensity at the focal zone and/or move the location of the focal zone.
For example, the speed of sound through fat is approximately 1460 meters per second (m/s), while the speed of sound through muscle is approximately 1600 meters per second (m/s). The speed of sound through bone tissue is much faster, for example, approximately 3000 meters per second (m/s) for skull bone tissue. The speed of sound also varies in different organs. For example, the speed of sound in brain tissue is approximately 1570 meters per second (m/s), approximately 1555 meters per second (m/s) in the liver, and approximately 1565 meters per second (m/s) in the kidney.
The relative phases (alternatively relative “time shift”) at which the transducer elements need to be driven to result in a focus at the target location depend on the relative location and orientation of the transducer surface and the target, as well as on the dimensions and acoustic material properties (e.g., sound velocities) of the tissue or tissues between them (i.e., the “target tissue”). Thus, to the extent the geometry and acoustic material properties are known, the relative phases (and, optionally, amplitudes) can be calculated, as described, for example, in U.S. Pat. Nos. 6,612,988, 6,770,031, and 7,344,509, the entire disclosures of which are hereby incorporated by reference. In practice, however, knowledge of these parameters is often too incomplete or imprecise to enable high-quality focusing based on computations of the relative phases alone. For example, when ultrasound is focused into the brain to treat a tumor, the skull in the acoustic path may cause aberrations that are not readily ascertainable. In such situations, treatment is typically preceded by an auto-focusing procedure in which, iteratively, an ultrasound focus is generated at or near the target, the quality of the focus is measured (using, e.g., thermal imaging or acoustic radiation force imaging (ARFI)), and experimental feedback is used to adjust the phases of the transducer elements to achieve sufficient focus quality.
The auto-focusing procedure may thus take a substantial amount of time, which may render it impracticable or, at the least, inconvenient for a patient. While the effect of pre-therapeutic sonications may be minimized by employing an imaging technique that requires only low acoustic intensity (e.g., ARFI), it is generally desirable to limit the number of sonications prior to treatment. Accordingly, there is a need for more efficient ways of focusing a phased array of transducer element to create a high-quality ultrasound focus.
Another common technique for focusing ultrasound is by using a shaped lens with an acoustic velocity (i.e. speed of sound) that differs from adjoining air, tissue, or material to bend acoustic waves. Most standard ultrasound focusing lenses employ a single concave lens. However, a single concave lens focusing system for ultrasound has limitations, including limitations. Ultrasound lenses comprised of a single concave lens are limited with regard to the range of focal lengths that can be achieved with a lens of a particular cross sectional area. Short focal lengths cannot be achieved with smaller cross sectional areas appropriate for systems affixed to the head or skull. Neuromodulation of superficial brain regions with an appropriate transcranial ultrasound system would be advantageous due to the importance of such superficial brain regions (e.g. cerebral cortex) to sensory, motor, higher cognitive function, and other brain functions.
To affect brain function transcranial ultrasound neuromodulation requires appropriate ultrasound waveform parameters, including acoustic frequencies generally less than about 10 MHz, spatial-peak temporal-average intensity generally less than about 10 W/cm2 (e.g., between 0.5 and 10 W/cm2), and appropriate pulsing and other waveform characteristics to ensure that heating of a targeted brain region does not exceed about 2 degrees Celsius for more than about 5 seconds. Transcranial ultrasound neuromodulation induces neuromodulation primarily through vibrational or mechanical mechanisms. Noninvasive and nondestructive transcranial ultrasound neuromodulation is in contrast to other transcranial ultrasound based techniques that use a combination of parameters to disrupt, damage, destroy, or otherwise affect neuronal cell populations so that they do not function properly and/or cause heating to damage or ablate tissue.
As by in the article Lindsey (Lindsey B D, Smith S W. Refraction Correction in 3D Transcranial Ultrasound Imaging. Ultrasonic imaging. 2014; 36(1):35-54. doi:10.1177/0161734613510287), Image quality in transcranial ultrasound remains limited by the deleterious effects of the skull, including attenuation, aberration, refraction, and mode conversion. Effects of attenuation may be reduced by positioning the probe within an acoustic window in the temporal bone; however, this window is absent in 8% to 29% of individuals. Transmitting with large, lower frequency (˜1 MHz) array probes may help reduce the dependence on acoustic window quality. The effects of aberration induced by spatially inhomogeneous layers having a different longitudinal wave velocity from that typically assumed by the ultrasound scanner (c=1540 m/s) may be addressed by one of the many techniques for phase aberration correction. Some of these aberration correction techniques include inherent correction for refraction using either ultrasound-based measurements in two-dimensional (2D) imaging or computed-tomography-based measurements in three-dimensional (3D) therapy though refraction correction in 3D imaging has not been demonstrated in vivo. Other techniques have modeled aberration as a distributed phenomenon rather than as a single spatially varying layer. Previously addressed anatomical sources of aberration include layers of bone in the skull (c≈2800 m/s, commonly within 15% variation due to difference in bone porosity and thickness) or layers of fat (c≈1450).
As discussed in the journal article by Ding et. al., Phys. Med. Biol. 60 (2015) 3975-3998, ultrasound penetration through the skull is better, with less energy deposition within the skull bone itself, at around 0.5 MHz compared with higher frequencies.
There is a need for more efficient ways of focusing a phased array of transducer element to create a high-quality ultrasound focus without recourse to imaging model of the brain by MRI or Ultrasound.
As illustrated in
Yet, as illustrated in
Since a beam of acoustic energy has a relatively wide aperture where it enters the body, different parts of the acoustic energy, such as 123 and 124, may pass through different intervening tissue layer thickness between, which may shift the effective relative time delay of acoustic energy transmitted from respective transducer elements upon arrival to the focal zone. This phase shifting may decrease the constructive interference of the acoustic energy at the focal zone, or may even move the focal zone in an unpredictable manner. For example, an intervening skull bone layer thickness difference of 1.5 mm may introduce a phase shift of 180° at an ultrasonic frequency of one Megahertz (1 MHz), which would change desired constructive interference at the focal zone into destructive interference.
In preferred embodiments, ultrasound frequency lower than 1 MHz, such as between 100 KHz and 500 KHz, is used in order to reduce the aberration effect of non-bone tissue inhomogeneity in the brain.
As illustrated in
The key problem is how to determine the adjustments set {dFn}. The present invention provides a method and system for determine the adjustments set {dFn} and thereby creating an improved focus when the focused ultrasound waves path pass through an intervening second-tissue intermediate-layer bounded by first and second boundary surfaces within otherwise approximately uniform medium 99 in the path from the transducer surface to the focus peak.
In general, as in prior art, we conceptualize the method procedure as composed of two major stages: (i) scanning procedure “PROC-1”, and (ii) adjusted irradiation application procedure “PROC-2”.
The scanning procedure PROC-1 can be sub-divided into two prominent sub-tasks: (a) physical scanning procedure “SCAN”, and (b) computational analysis “ANALYSIS” from which the key outcome is the determination of the adjustments set {dFn}.
The adjusted irradiation application procedure “PROC-2” can be sub-divided into two prominent sub-tasks: (a) Input parameters set-up, and (b) irradiation application process.
The adjusted irradiation application procedure “PROC-2” is similar to prior art. Given the adjustments set {dFn} from PROC-1, the system is activating the emitters array with the corrected set {F′n} of input parameters phases. The irradiation application process itself is determined by clinical goals (e.g., nerve stimulation, tissue ablation, etc. . . . ).
The innovation is primarily contained in the preparatory scanning process PROC-1 method and apparatus and from it the core outcome is the values of the adjustments set {dFn} of input parameters that is used to define the corrected set {F′n} of phases.
As known in prior art, there are multiple methods of determining the adjustments set {dFn} if there is a given geometry of the intermediate layer (e.g., skull bone layer). The problem is how to determine the skull properties and geometry non-invasively. For concreteness we define the intermediate layer geometry by knowledge of first boundary surface 151 shape function B1(x,y,z) and the second layer 152 shape function B2(x,y,z).
In prior art, for skull bone, the skull bone layer geometry is determined from MRI or CT imaging. Thus, in prior art the SCAN sub-task scanning procedure PROC-1 is done by MRI or CT scanning, from which various skull parameters are determined by supplemental external information. e.g., from multiple slices of MRI images a full 3D skull bone section shape is reconstructed. In addition, external information concerning speed of sound in the bone is supplemented to predict and determine the supposed time shifts created by the skull bone on ultrasound. i.e., the MRI or CT scanning is NOT by itself directly measuring ultrasound phase shift or time shift due to passage through the skull layer.
In prior art ANALYSIS sub-task, each phase shift correction dFn to be applied to individual emitter En is determined by going through a geometrical reconstruction of the intermediate layer shape and considering the particular path of the ultrasound from the emitter En to the intended focus location.
In contrast, the present invention: (i) uses ultrasound emitters array not only for the irradiation PROC-2 procedure, but also for the scanning PROC-1 procedure, thereby eliminating completely the need for non-ultrasound MRI or CT at any step of the full procedure (PROC-1 and PROC-2) method and apparatus; (ii) the ultrasound phase shift or time shift due to passage through the skull layer is directly measured by the SCAN procedure ultrasound scan, and (iii) the SCAN procedure ultrasound scan method in the present invention is different from what is commonly understood as “ultrasound scanning” in prior art. In particular, while the majority of the array elements {En} are activated simultaneously during the irradiation PROC-2 procedure (as in common ultrasound scanning), only individual emitter elements En (or a small fraction of the array elements, e.g., less than 20%) are activated simultaneously during the SCAN process of PROC-1.
In prior art scanning with ultrasound phase arrays, what is conventionally understood to be ultrasound scanning: (i) the array elements are operated to radiate simultaneously, (ii) a focus beam is used, and (iii) target area scanning is performed by steering the beam focus by way of modifying the relative phase shifts between simultaneously activated array elements. In order to find and trace the skull outer surface, prior art phased array scan is moving the focus over a volume within which the skull is assumed to be residing somewhere. In contrast, non of the above is conducted in the present invention ultrasound SCAN procedure.
The present invention SCAN sub-task is characterized by that: (i) the array ultrasound elements are operated serially in time, such that individual array element (or small groups of elements the majority of which consisting of less than 10% of the number array elements) are operated on after the other, and preferably after the previous element signal reflection have been measured; (ii) the ultrasound beam is not focused; (iii) target area scanning is performed by NOT by steering a beam focus, but instead by way of each individual array element (or small group of elements) measuring the small section of the intermediate layer (e.g., skull bone) closest to it.
In addition, in preferred embodiments of the present invention ANALYSIS sub-task, each phase shift correction dFn to be applied to individual emitter En is determined directly from the associated individual emitter En SCAN step, without going through a geometrical reconstruction of the intermediate layer shape and without considering the particular path of the ultrasound from the emitter to the intended focus.
In preferred embodiments,
dFn=(V2/V1−1)*Tn/2
i.e.,
F′(n)=F0(n)+(V2/V1−1)*Tn/2,
where V1 is an assumed average speed of sound in the interior brain tissue (preferably within 10% accuracy), and V2 is an assumed average speed of sound in the intermediate layer (e.g., skull bone), preferably within 20% accuracy (or better). The time Tn is determined from the time difference between the reflected signals 171 from the intermediate layer first surface 151 and the reflected signals 172 from the intermediate layer second surface 152.
In other preferred embodiments, ANALYSIS sub-task is more conventionally performed, such that each phase shift correction dFn to be applied to individual emitter En is computed by going through a geometrical reconstruction of the intermediate layer shape and considering the particular path of the ultrasound from the emitter En to the intended focus location. Yet, in the present invention the geometrical reconstruction of the intermediate layer shape and thickness are determined in a novel way.
For example, as illustrated in
We interpret and associate each measured Tn as an estimation of the local time difference of ultrasound to so twice across the local skull bone thickness nearest to the emitter element En.
In preferred embodiments, the adjustments elements dFn forming the adjustment set {dFn} are each a function of Tn, V1 and V2. Preferably, as noted before,
dFn=(V2/V1−1)*Tn/2.
We interpret and associate an estimated local skull bone thickness Wn, where
Wn=V2*Tn/2,
as the local skull bone thickness nearest to the emitter element En. V2, the speed of sound in the skull bone, is not necessarily uniform across the skull. For example, it may vary between thicker and thinner areas of the skull bone.
In preferred embodiments, the first estimation of Wn is given with a pre-determined selected V2 value (e.g., V2=2800 m/s, or 2600 m/s). Then, an iterated estimation of Wn is adjusted based on the value of first estimation of Wn. For example, is Wn is corrected with assuming V2 higher in thinner skull areas and lower at thicker skull areas.
In addition to the phase corrections, amplitude corrections can be determined from the reflected test signals from individual emitters.
In addition, to maximize the transmitted intensity, a step of frequency-test scan is added to be performed prior to treatment application PROC-2, or prior to the SCAN procedure. In the frequency-test scan, the goal is to find the frequency of maximum transmission through the intermediate layer in order to minimize loss of intensity at the focus (due to reflection from the intermediate bone layer) and in order to minimize heat deposition within the intermediate bone layer. In a frequency-scan, for a local region the reflected signal intensity is measured while changing the emitter activation frequency range around a central work frequency, e.g., within 10% deviation from the central work frequency. For example, if the work frequency is chosen to be 500 KHz, a scan of frequency range between 450 KHz and 550 KHz is performed. Minimum local of reflected signal intensity indicates maximum local transmission at the associated frequency. In preferred embodiments, the frequency scan is performed for the array activation as a whole (rather than for local regions).
A presently preferred embodiment of the invention will be described in detail, in conjunction with the accompanying drawings, in which:
The invention is herein described, by way of example only, with reference to the accompanying drawings. With specific reference now to the drawings in detail, it is stressed that the particulars shown are by way of example and for purposes of illustrative discussion of the preferred embodiments of the exemplary system only and are presented in the cause of providing what is believed to be a useful and readily understood description of the principles and conceptual aspects of the invention. In this regard, no attempt is made to show structural details of the invention in more detail than is necessary for a fundamental understanding of the invention, the description taken with the drawings making apparent to those skilled in the art how several forms of the invention may be embodied in practice and how to make and use the embodiments.
For brevity, some explicit combinations of various features are not explicitly illustrated in the figures and/or described. It is now disclosed that any combination of the method or device features disclosed herein can be combined in any manner—including any combination of features—any combination of features can be included in any embodiment and/or omitted from any embodiments.
For ease of reference the following numbers in the figures are meant to refer to as follows
It helps to highlight from the outset certain distinguishing feature of the present invention preferred embodiments in comparison with prior art.
As illustrated in
Since any of the above and also combination of the above transducer array models can be used to realize the present invention, we use a generic representation, illustrated in
As illustrated in
As illustrated in
The key problem is how to determine the adjustments set {dFn}. The present invention provides a method and system for determine the adjustments set {dFn} and thereby creating an improved focus when the focused ultrasound waves path pass through an intervening second-tissue intermediate-layer bounded by first and second boundary surfaces within otherwise approximately uniform medium 99 in the path from the transducer surface to the focus peak.
We conceptualize the method as a procedure composed of two major stages: (i) scanning procedure “PROC-1”, and (ii) adjusted irradiation application procedure “PROC-2”.
The scanning procedure PROC-1 can be sub-divided into two prominent sub-tasks: (a) physical scanning procedure “SCAN”, and (b) computational analysis “ANALYSIS” from which the key outcome is the determination of a time-set {Tn} from which is determined the adjustments set {dFn}, where each element dFn is a function of corresponding Tn.
The adjusted irradiation application procedure “PROC-2” can be sub-divided into two prominent sub-tasks: (a) Input parameters set-up, and (b) irradiation application process.
In the adjusted irradiation application procedure “PROC-2”, the adjustments set {dFn} from PROC-1, the system is activating the emitters array with the corrected set {F′n} of input parameters phases. The irradiation application process itself is determined by clinical goals (e.g., nerve stimulation, tissue ablation, etc. . . . ).
The innovation is primarily contained in the preparatory scanning process PROC-1 method and apparatus and from it the core outcome is the values of the adjustments set {dFn} of input parameters that is used to define the corrected set {F′n} of phases.
As illustrated in
In the path to the focus peak 111 from the emitter 141n, the sound wave passes once through the intermediate layer 150 width. In contrast, the reflected test signals, 172 and 174, from boundary 152, each passes twice through the intermediate layer 150 width (corresponding to the incident and reflected portions of the path within the layer 150). Therefore, half of the reflected test signals time difference Tn, i.e., (Tn/2), is a good approximation to the time shift contribution of the intermediate layer 150 to the total travel time of the focused beam from the particular emitter 141n to the focus peak location 111. Therefore, in preferred embodiments of the present invention, adjusting the relative time shift by subtracting (Tn/2) from F0(n) substantially eliminates the phase shift contribution of the intermediate layer 150 to the total travel path from the emitter 141n to the focus peak 111. Hence, in preferred embodiments the adjusted delay set {F′(n)}={(F0(n)+dFn)} is then use for irradiating the target tissue using the ultrasound transducer array emitter elements {E(n)} having a corrected-delay {F(n)}={F′(n)} to create a focus peak 131 with an adjusted focal zone 139 of smaller cross section area than if irradiated with unadjusted delay set {F0(n)}.
It remains to be decided which receiver sensor measurement to use for the determination of Tn. Ideally, one would like to have a reflected path which is exactly twice the length of the focused beam path within the intermediate layer 150.
One preferred approximation, as illustrated for the path 172, is a test signal path for which at least the incident portion of the test signal path to be the same as the focused beam path 170 to the focus peak 111. For such an embodiment, fast switching needs to be operated to switch the emitter transducer to receiver mode of operation within the duration of the reflection time. To a good approximation, for ray 370, defined to be the ultrasound propagation “ray” from emitter 341 as would be in the selected uniform medium 99 traversing the focus peak 111, the path of ray 370 through the intermediate layer 150 is very close to the real path of ultrasound ray from emitter 141n to the focus peak 111.
Another preferred embodiment approximation is one for which the same physical transducer array element is used for both test signal emitter and as receiver sensor. Another preferred embodiment approximation is one for which, as illustrated for the path 171, fora given emitter 141n the associated receiver sensor 161n is a neighboring (e.g., nearest neighbor) transducer sensor element 161n. A question in such embodiments is which direction of neighbor to choose as sensor placement relative to the emitter element. For example, should it be one to the left or to the right of the emitter element. The better approximation depends on the relative curvature of the entry boundary 151 of the intermediate layer 150 compared with the curvature of the focus beam at that boundary surface. For example, as illustrated in
There is no need for unique exclusive one-to-one association of a sensor to an emitter. In some preferred embodiments, the number receiver sensors is smaller than the number of emitter elements in {E(n)}. For example, as illustrated in
In some preferred embodiments, the emitter and receiver sets are physically distinct. In some other preferred embodiments, the emitter and receiver sets are overlapping. For example, referring to
It is preferred to be able to selectively activate individual emitter and/or receiver elements from the physical arrays sets. For example, for test signal and adjustment procedures, one is preferably activating the array emitter elements serially or in sub-sections (i.e., not all together as for focus creation), and respectively receiving and/or analyzing the echo received signals only of the associated receiver sensors. That is relevant both for embodiments where the emitter and receiver sets are physically distinct and for embodiments where they are overlapping. In preferred embodiments, the switching control and activation of the emitter 112 and/or receiver 116 arrays are managed by a controller module 255.
A focusing transducer 110 comprising an emitter phased array {E(n)} 112 of ultrasound transducer elements for generating an ultrasound focus in the target tissue. At least most transducer elements having means connected thereto for variably setting a delay for that emitter transducer E(n), including a delay module for setting delay set {F(n)} of signals delay to associated emitter elements {E(n)}.
A receiver array 116 comprising a plurality of ultrasound receiver elements {R(n)} associated with emitter elements {E(n)}; The receiver array 116 is connected to and controlled by a receiver control module 170.
A transmit/receive controller “T/R controller” module 255 for directing the activation and switching of individual elements of the emitter and receiver arrays. The T/R module 255 comprising connections to the emitters array 112, to the receiver array 116, to the signal generator module 140 and to the control & computation module 160.
An emitter mux module 145 receives input signal from the signal generator 140, the associated activation delay set {Tn} from control module 160, and the array elements activation directed by T/R module 255. The emitter mux 145 transmit the activation signals to the emitter array 112 elements {E(n)}.
For focusing-mode all, or at least the majority, of the emitter array 112 elements {E(n)} are activated simultaneously, with associated delay set {F(n)} of a common action-signal, to generate a focus peak at a certain location.
For adjustment-mode a minority, preferably one, of the emitter elements E(n) is activated at a time, preferably with a distinct “test-signal”, which may be preferably different from the action-signal. The test signal reflected echo-signal is received at an associated receiver element of the receiver set 116 and transmitted for analysis to receiver control module 170. It is expected that the received reflected signal would include multiple reflections from various material boundaries such as skin surface, fatty tissue, and the intermediate skull bone layer 150. As highlighted by the table of
After adjustment mode procedure, the focusing mode is activated with the emitter array 112 driven with an adjusted time-delay set {F′(n)}={F0(n)+dFn} to create an adjusted focus peak.
For effective focus improvement by the invention adjustment procedure. The positioning module 165 maintains the transducer module 110 at a fixed orientation relative to the target tissue for both the adjustment procedure and the focusing mode activation.
As previously discussed, better focusing adjustment would be obtained if the test signals path through the intermediate tissue layer is better matching (i.e., closer) to the path through the intermediate layer of focusing beam to the focus peak location. As illustrated in
In preferred embodiments, the preferred associated test-receiver for optimal path matching is not fixed for all intermediate layers, but is dependent on the orientation of the intermediate tissue boundary layers relative to the transducer focusing-axis 310. For example, as illustrated in
As illustrated in
The selection of receiver sensors at which time of reflected beams is determining the time delay correction adjustment.
The time difference between these reflection signals T2(n)−T1(n) is determined by the time-shift determiner module 172 to create the reflection time difference Tn=T2(n)−T1(n). Repeating the process serially for multiple, preferably most or preferably all, of the emitter elements, lead to obtaining a reflection time difference set {Tn}.
After adjustment mode procedure, the focusing mode is activated with the emitter array 112 driven with an adjusted time-delay set {F′(n)}={F0(n)+dFn} to create an adjusted focus peak.
The selection of receiver sensors at which time of reflected beams is determining the time delay correction adjustment.
The previously discussed embodiment, with reference to
In preferred embodiments, the sensor for timing T1 of the arrival of reflection 571 is timed at receiver 561 at which reflected beam 571 is detected at peak intensity. In preferred embodiments, the sensor for timing T2 the arrival of reflection 572 is timed at receiver 562 at which reflected beam 572 is detected at peak intensity. Repeating the test process for each emitter E(n) we obtain the reflected signals arrival times T1(n) and T2(n).
The time difference between these reflection signals T2(n)−T1(n) is determined by the time-shift determiner module 172 to create the reflection time difference Tn=T2(n)−T1(n). Repeating the process serially for multiple, preferably most or preferably all, of the emitter elements, lead to obtaining a reflection time difference set {Tn}.
After adjustment mode procedure, the focusing mode is activated with the emitter array 112 driven with an adjusted time-delay set {F′(n)}={F0(n)+dFn} to create an adjusted focus peak.
In addition to phase corrections (i.e., time delays) for individual emitters En, also amplitude corrections are determined in preferred embodiments. For example, for a given emitter test signal, the fractional intensity which is collected at the receiver peak (and optionally including the intensity of the nearest neighbors receivers also) for both the first boundary reflection and second boundary layer reflected echo signal is indicative of the remaining transmitted intensity that reaches the focus peak. Higher reflected fraction RI(n) means lower transmitted intensity fraction “TI(n)” contributing to the focus peak. Hence, in preferred embodiments, the amplitude correction A(n) for each emitter En of the set {En} is a function of TI(n), e.g., proportional to 1/TI(n). For example, in preferred embodiments, in order to obtain more uniform contribution to the focus peak from each emitter En, the intensity emitted from emitter En is set to be [1/TI(n)]*A0, where A0 if the intended average intensity of the transducer emitters array.
In addition, to maximize the transmitted intensity, a step of frequency-test scan is performed prior to treatment application. Typical skull bone thickness ranges between 4 mm to 12 mm. Typical ultrasound treatment frequency ranges between 0.25 MHz to 2 MHz, which at average speed of sound in bone of 3000 m/s translates to a range of wavelength between 1.5 mm to 12 mm. Hence, ¼ wavelengths range in size between 0.375 mm to 3 mm. Due to the ¼ wavelength maximum transmission effect, for a given skull bone sample, for any given desired central frequency W0 for treatment (e.g., 1 MHz, with associated wavelength of 3 mm for bone speed of sound of 3000 m/s) there will be a particular optimal frequency W1 in the neighborhood of W0 for which the transmitted ultrasound total beam intensity across the skull bone layer is maximized. This can respectively be detected as a minimum in the reflected intensity.
Since the intermediate layer (e.g., skull bone) thickness is non-uniform, it is possible and even likely that, the optimal frequency W1 of maximum transmitted intensity is different for different sub-regions of the surface area 155 under the full transducer area. Therefore, in preferred embodiments, as illustrated in
The optimal frequency W1 is determined independently for each sub-section. Then, the treatment is delivered with each sub-section driven at its own local optimal frequency, in parallel or serially with other sub-sections of the transducer array.
In all of the above examples, the key in the preparatory SCAN process is to get separate measurements of individual emitters En (or small local group of emitters). In practice, the signals from sufficiently distant emitters do not mix or interfere at the small distance of the SCAN reflection measurements. Therefore, in preferred embodiments, the SCAN process can be performed simultaneously on multiple emitters sub-set of the full array set {En} such that the distance between the array elements is larger than the distance between each array element and the intermediate skull layer 151.
In preferred embodiments of the present invention the SCAN procedure is the array 112 is positioned close to the layer surface 151. In preferred embodiments of the present invention the curvature of the array is very close to concentric with the curvature of the intermediate layer, typically between 6 cm and 12 cm radius of curvature. Since different areas of the skull have some difference in curvature, and also there is a range of curvature differences between human individuals, the array used in a particular instance can be change-matched to the area of the skull that is irradiated in practice.
The SCAN procedure in the simulation consisted of serially performing the same SCAN step on all the 20 emitters forming the emitters array 112 set {En}.
Finally,
In the above preferred embodiments examples the PROC-1 SCAN process and the PROC-2 irradiation process where both performed with the emitters array 112 at the same relative position with respect to the intermediate layer 150. But this is not necessarily the case.
In preferred embodiments, in order to enable preferred embodiment where the PROC-1 SCAN process and the PROC-2 irradiation process where performed with the emitters array 112 at the different relative position with respect to the intermediate layer 150, we introduce in PROC-1 an optional supplemental ANALYSIS step—labeled LAYER RECONSTRUCTION—and introduce in PROC-2 a supplemental irradiation preparation optional step—labeled PATH ANALYSIS. The key additional step is the LAYER RECONSTRUCTION sub-process at the end of the ANALYSIS process.
The LAYER RECONSTRUCTION outcome is a geometrical reconstruction of a portion of the intermediate layer which is facing the emitters array 112 during the SCAN process. For concreteness, we define the intermediate layer geometry by estimation of outer boundary surface 151 shape function B1(x,y,z) and the inner layer 152 shape function B2(x,y,z). In the prior-art, these were deduced from external scans by MRI or CT. In contrast, in the present invention the geometrical functions reconstruction is derived from the set of parameters obtained in the invention SCAN process and ANALYSIS. As indicated in
As known in prior art, there are multiple methods of determining the adjustments set {dFn} if there is a given geometry of the intermediate layer (e.g., skull bone layer). Hence, there is no need for us to elaborate how the irradiation process is done once the LAYER RECONSTRUCTION outcome is given.
The usage of these input parameters to reconstruct the intermediate layer outer boundary function B1 and inner boundary function B2 is a matter of approximation. Any approximation procedure gives a mathematically non-identical outcome. Moreover, small deviation from precise functional reconstruction of the true intermediate layer geometry may be of little significance to the intended clinical outcome of the irradiation procedure. Therefore, it is not the particular functional approximation which is the essence of the inventions, but it is the set of above noted input parameters obtained from the SCAN process and ANALYSIS. To highlight this statement, the illustrated approximation procedure in
The table illustrated in
All publications, patent applications, patents, and other references mentioned herein are incorporated by reference in their entirety. In case of conflict, the patent specification, including definitions, will prevail. In addition, the materials, methods, and examples are illustrative only and not intended to be limiting.
The present invention has been described using detailed descriptions of embodiments thereof that are provided by way of example and are not intended to limit the scope of the invention. The described embodiments comprise different features, not all of which are required in all embodiments of the invention. Some embodiments of the present invention utilize only some of the features or possible combinations of the features. Variations of embodiments of the present invention that are described and embodiments of the present invention comprising different combinations of features noted in the described embodiments will occur to persons of the art.
Number | Date | Country | |
---|---|---|---|
62462398 | Feb 2017 | US | |
62468473 | Mar 2017 | US | |
62474715 | Mar 2017 | US | |
62492193 | Apr 2017 | US |