This invention relates to a method for determining, at any arbitrary location within a target region of a soft solid, a local intensity of an acoustic field propagating in this target region. This invention also relates to an apparatus for implementing such a method.
In the meaning of the present invention, a soft solid is an organic tissue which can have an animal or vegetal origin. For instance, such a soft solid can be an organ of a human or animal body, or a part of a vegetable. A soft solid can also be a food element, e.g. cheese, a cosmetic compound, or a non-metallic part of an equipment such as prosthesis, this non-metallic part being made of a natural or synthetic material.
When an acoustic field propagates in a soft solid, in order to measure a physical parameter of this solid or in order to apply a treatment, it is essential that the spatial distribution of the acoustic field in the soft solid is precisely controlled, in particular in order not to damage the soft solid, e.g. by overheating.
For instance, diagnostic and therapeutic ultrasound applications rely on delivering a specific intensity distribution of an acoustic field over a given duration. In medical applications, the ultrasound beam may be unexpectedly attenuated, or deflected, or both, because of tissue heterogeneity. Ideally, beam alteration should be detected and corrected. But, as of today, there is no way to measure intensity in biological tissues. Beam alteration is therefore likely to go unnoticed, with potential consequences on safety and efficacy of the diagnostic and/or of the treatment.
The intensity distribution of an acoustic field in a region of interest can be estimated using an inverse problem approach. Specifically, in a first iteration, some initial guess of the intensity distribution is entered in the simulation. The difference between the field predicted from simulation and a field measured experimentally is calculated. In subsequent iterations, the intensity distribution entered in the simulation is varied and the difference between the field predicted from simulation and the field measured experimentally is minimized iteratively. An algorithm selects the intensity distribution that minimizes the discrepancy between simulations and experiment. However, the chosen solution may correspond to a local minimum. It may therefore be inaccurate. This approach may therefore be unreliable.
Using a viscoelastic model is known in the field of ultrasound imaging, as explained by Bercoff et al. in “The role of viscosity in the impulse diffraction field of elastic waves induced by the acoustic radiation force” (IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, Vol. 51, No. 11, November 2004). This document focuses on the determination of soft tissue elasticity and is not concerned with the actual intensity of an acoustic field applied for an experiment.
In “Monitored steady-state excitation and recovery (MSSER) radiation force imaging using viscoelastic models” (IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, Vol. 55, No. 7, July 2008) and in US-A-20100138163, Mauldin et at consider a steady-state situation for measuring an elasticity. Here again, this does not allow monitoring the intensity of an acoustic field used in an experiment.
US-A-2012/226158 discloses a system and a method for correcting errors in shear wave measurements. This method disclosed does not allow for the measurement of an acoustic intensity. This method is therefore fundamentally different from the method described below.
Usually, the power of the beam of the acoustic field in the target region, which depends on the ultra-sound beam distribution, is estimated by simulation. This simulation is based on a situation in a homogenous medium, such as water. This can be satisfactory when the soft solid is also homogeneous.
In practice, many soft solids, in particular biological tissues, are heterogeneous. They can include localized variations, such as tumors and/or fat with a distribution which depends, amongst others, on the size of the organism to be studied, for instance a patient. Thus, in real life applications of ultrasonic beams, the beam can be deflected, attenuated or dilated/diverging due to heterogeneities, so that the practitioner does not know how effective the diagnostic or the treatment is and for how long is should be applied.
In other words, the intensity of the ultrasonic field in the target region of the diagnostic and/or of the treatment is not controlled. One relies on the simulation made in water, which is not appropriate for the soft solid.
The same problem occurs when an acoustic field is used for non-medical purposes, e.g. for quality control in the food, cosmetic or mechanical industry, where heterogeneities may result from holes or fat in a sample.
The invention aims at solving these problems by proposing an improved method which actually allows determining a local intensity of an acoustic field propagating in a target region of a soft solid, thus allows controlling that such an acoustic field is being efficiently used.
To this end, the invention concerns a method for determining a local intensity of an acoustic field induced by an ultrasound source and propagating in a target region of a soft solid, at a position located within said target region, characterized in that said method includes at least the following steps:
In the meaning of the invention, “steady-state” means that the displacement or displacement difference measured at step d) is obtained in a steady-state steady state condition, in other words that it is invariant in time during the implementation of the method of the invention.
The building of the viscoelastic model can be performed with a Green's function or with any other suitable tool, such as the finite differences method, or using the finite elements method, as described by Eskandari et al. in “Viscoelastic characterization of soft tissue from dynamic finite element models” (Physics in medicine and biology 2008; 53(22): 6569-90).
In physics, an inverse problem is the process of determining the causes of a phenomenon, based on the observation of its effects. For simple phenomena, the solution can be as simple as solving an equation. On the other hand, complex phenomena require extensive numerical calculations performed using a computer or an array of computers. Methods for solving inverse problems are described in the book “Parameter Estimation and Inverse Problems”, (2018), Third Edition, Elsevier. ISBN 9780128134238, ISBN 9780128134238″ by Aster Richard; Borchers Brian, and Thurber Clifford. In the present invention, inversion of the viscoelastic model can be achieved using any of the known methods described in this book.
Thanks to the invention, it is possible to determine how efficient an acoustic beam is, in particular to take into account a possible alteration of the acoustic field, such as a deflection or a diverging configuration of the beam. When detected, such an alteration can be corrected, which improves the efficiency of the control or treatment performed with the acoustic field.
According to further aspects of the invention which are advantageous but not compulsory, the method might incorporate one or several of the following features, taken in any technically admissible configuration:
The invention also concerns an apparatus for implementing a method as mentioned here-above, characterized in that said apparatus comprises:
Such an apparatus is efficiently tailored to provide the same advantages as the method of the invention.
Advantageously, the electronic computation unit for building the viscoelastic model and the electronic computation unit for computing the amplitude of the intensity of the acoustic field are formed by a single piece of equipment.
Finally, the invention also concerns a therapeutic ultrasound process including steering of an ultrasound field towards a target region of a biological tissue, wherein the local intensity of the acoustic field is determined with a method as mentioned here-above.
The invention will be better understood on the basis of the following description which is given in correspondence with the appended figures and as an illustrative example, without restricting the object of the invention. In the annexed figures:
The apparatus A represented of
S6 represents a signal sent by ECU 6 to array 2 in order to control this acoustic source.
A soft solid 8, for instance a biological tissue, is installed within volume V4 and a target region 82 of this soft solid is subjected to the acoustic beam B originating from acoustic source 2.
Alternatively, if soft solid can stand alone, box 4 can be omitted.
In this example, soft solid is a part of a muscle of an animal to be studied.
According to a known feature of acoustic emitters/receptors, acoustic array 2 is also configured as a scanner for detecting an image of soft solid 8, this image showing the location of several positions of this soft solid within target region 82.
S2 represents a signal sent by ultrasonic array 2 to ECU 6, this signal S2 being representative of the scanned image detected.
Alternatively, two different pieces of equipment can be used instead of array 2, that is one for emitting beam B and another one for scanning the image of target region 82.
ECU 6 includes some electronic components, such as one or several microprocessors and one or several memories, and some computer programs to automatically implement the method of the invention.
When beam B is applied to soft solid 8, it is essential to know the effect of this beam within this soft solid. This effect depends on the local intensity of the acoustic field AF. This local intensity can be determined with the method of the invention.
Let u denote a displacement observed at a given position within target region 82 and {right arrow over (i)} denote the instantaneous acoustic field intensity vector, or Poynting vector, at the same position.
First Method of the Invention
The method of the invention relies upon the fact that the acoustic field intensity {right arrow over (i)}, at a given position within target region 82, can be considered as proportional to a force F applied to this position. On the other hand, the displacement u of the soft solid at this position can also be considered as proportional to a weighted sum of forces. Thus, a proportionality relationship exists between the acoustic field intensity {right arrow over (i)} and the displacement u obtained at the same position.
A viscoelastic model M, such as the one represented on
Note here that the viscoelastic model M is not necessarily linear. In other words, the curve of
The method of the invention also relies on the fact that, if the displacement u and the viscoelastic model M are known, an inversion of the viscoelastic model M allows determination of the acoustic field intensity {right arrow over (i)}.
On the other hand, according to the literature, in particular Bercoff et al. cited here-above, the viscoelastic model M depends, amongst others, on the following three parameters:
A first method according to the invention is schematically represented on
Determination of Parameters α, μ and c
In a first step 100, the method begins and some memories of ECU 6 are emptied or reset.
In a subsequent step 102, the ultrasound attenuation coefficient α is determined, either by access to the literature or experimentally, as taught by Ophir et al in Ophir J, in. “Attenuation estimation in reflection: progress and prospects”. (Ultrason Imaging. 1984 October; 6(4):349-95. Review), whose content is incorporated herein by reference.
In another step 104, the shear modulus μ is determined, either by access to the literature or experimentally, as taught by Muthupillai et al in “Magnetic resonance elastography by direct visualization of propagating acoustic strain waves” (Science. 1995 Sep. 29; 269(5232):1854-7) or by Sarvazyan et al in “Shear wave elasticity imaging: a new ultrasonic technology of medical diagnostics” (Ultrasound Med Biol. 1998 November; 24(9):1419-35) whose content is incorporated herein by reference.
At step 104, experimental estimation of the shear modulus μ involves
Estimation of the shear modulus is not possible in the region where the shear wave is induced. Hence, in step i mentioned here above, the shear wave is initiated in the vicinity of the target region of interest, but not exactly at the region of interest.
In another step 106, the speed of sound c is derived from the literature. Alternatively, it can be measured.
Steps 102, 104 and 106 can be simultaneous or occur in any order. Each of these steps includes an operation of storing, in a memory of ECU 6, the value of the parameter detected.
Building of a Viscoelastic Model
In order to be able to actually compute a viscoelastic model and invert it at a later stage one uses assumptions allowing simplification.
The first assumption is that the acoustic field AF whose intensity {right arrow over (i)} is to be determined is steady invariant in time. In other words, this acoustic field is supposed to have a constant intensity, even if this field induces vibration over time of some particles in the target region 82.
The second assumption is that the acoustic field is applied for a duration such that the acoustic field induces a steady-state localized deformation of the soft solid in the target region.
Let {right arrow over (i)}({right arrow over (r)}, t) represent the instantaneous acoustic intensity vector, or Poynting vector, at a given position {right arrow over (r)} and time t. Assume that the acoustic wave is cyclic, with repetition time T. Let {right arrow over (I)}({right arrow over (r)}, t) represent the time averaged acoustic intensity vector. Let I({right arrow over (r)}, t) be the amplitude of the time-averaged intensity vector.
By definition, the following equation prevails:
The radiation force distribution {right arrow over (f)}({right arrow over (r)}, t) induced by the acoustic field is given by Torr et al in “The acoustic radiation force” (American Journal of Physics 1984; 52(5):402-8) whose content is incorporated by reference. The following equation prevails:
where α is the attenuation coefficient of sound and c is the speed of sound identified here above.
The displacement distribution {right arrow over (u)}({right arrow over (r)}, t) induced by the radiation force in a homogeneous isotropic viscoelastic medium is given by Bercoff et al in “The role of viscosity in the impulse diffraction field of elastic waves induced by the acoustic radiation force” (IEEE Trans Ultrason Ferroelectr Freq Control. 2004 November; 51(11):1523-36), whose content is incorporated by reference. The following equation prevails:
ui({right arrow over (r)},t)=Σj=13([fj{circle around (×)}{right arrow over (r)},tgij]({right arrow over (r)},t)) (3)
where
The Green's function also depends on parameters μ, η, α and c, where η is the shear viscosity in the soft solid 8. Hence the symbol {right arrow over (g)}({right arrow over (r)},t) is actually a shorthand notation for {right arrow over (g)}({right arrow over (r)},t,α,c,μ,η).
In the most general case, the displacement distribution induced by the instantaneous acoustic intensity can be predicted in some region of interest, by combining equations (2) and (3).
In most applications, the acoustic beam B exhibits a specific pattern, or shape, in space. Moreover, its intensity is modulated to follow a specific waveform in time. Shaping in space is typically achieved electronically using focusing or “beamforming” techniques for driving ultrasonic source 2. Modulation in time is typically achieved by driving the setup with specific acoustic power waveforms. For example, the acoustic power waveform may be a rectangle function, resulting in constant power being applied over a specific duration. Assuming that the shape of the acoustic field is time-invariant, as mentioned in the first assumption here above, the intensity distribution can be as mentioned in the following equation:
{right arrow over (I)}({right arrow over (r)},t)=I0·W(t)·{right arrow over (S)}({right arrow over (r)}) (4)
where I0 is the peak value of local intensity, {right arrow over (r0)} is the position where this peak is expected to occur, the vector field {right arrow over (S)}({right arrow over (r)}) represents the shape of the beam, i.e. the spatial distribution of acoustic intensity with amplitude normalized by I0, and scalar function W(t) represents the temporal waveform of acoustic intensity, also normalized by I0. I0. can also be considered as the amplitude of the intensity of the acoustic field at an arbitrary position {right arrow over (r0)}
Under these assumptions, equation (3) becomes:
where Sj({right arrow over (r)}) is the component in the xj direction of the normalized intensity field at the given position {right arrow over (r)}, in other words the shape of the acoustic field.
Equation (5) shows that the entire displacement distribution ui({right arrow over (r)},t) is proportional to peak intensity I0. As a consequence, the experimental value of peak intensity I0 can be estimated in three steps, namely:
The position {right arrow over (r1)} is preferably chosen to be at the location {right arrow over (r0)} of the intensity peak I0, or in its vicinity, so that ui,exp({right arrow over (r1)}) is maximum and can be determined with high precision.
The computations leading to equation (6) constitute an inversion of the viscoelastic model M represented of
Since the spatial distribution of the intensity {right arrow over (i)} is as shown in equation (4) and since functions W(t) and {right arrow over (S)}({right arrow over (r)}) are known, it is enough to know Io,exp in order to know {right arrow over (I)}({right arrow over (r)},t).
A first limitation of the technique described above is that shear viscosity i is usually unknown in biological tissues. A solution can be found using steady-state displacement, as mentioned here-above with the second assumption. During steady-state, viscous forces are nulled, and the resulting displacement is independent of shear viscosity.
Steady-state condition is achieved using long excitation with constant power for exiting the ultrasound array 2. In this example, the corresponding excitation waveform W(t) may be a Heaviside function:
The component in the xi direction of the steady-state displacement is given by:
where
is the steady-state Green's function of the medium, i.e. the displacement field induced by a Heaviside excitation force applied at a position source. Equation (8) is obtained by plugging equation (7) into equation (5), then writing the convolution product in integral form, and replacing W(t) by 1 for t>0.
Thus, the displacement u induced at one or several position(s) of interest {right arrow over (r)} by the acoustic field may be pre-calculated for any intensity value I0, using equation (8) and the values α, and μ and c determined in steps 102, 104 and 106. This allows creating, in a further step 108, a displacement image DI of the soft solid in target region 82, by grouping the displacement values of the different positions of interest.
It is possible to work with one position of interest {right arrow over (r)} only. In such a case, the displacement image DI includes one pixel only, which corresponds to the displacement u pre-calculated for this position of interest.
The displacement images DI, that is the displacement values, obtained at steps 108 are then used in a subsequent step 110, in order to build a viscoelastic model M, which is a relationship between the local intensity I0 of the acoustic beam B and the induced displacement u. For each direction (i=1, 2 or 3) of the coordinate system, this model can take the shape represented by the curve of
In the most general case, and especially when high acoustic intensity is being used, nonlinear wave propagation may occur. The relationship between displacement and intensity can take any shape, such as the one illustrated in
In order to build the calibration chart, step 108 is repeated for different intensities I0 of acoustic beam B.
However in the specific case of low acoustic intensity, acoustic wave propagation can be assumed to be linear. The relationship between displacement and intensity is a straight line, as illustrated in
Alternatively, the model can be an elastic model, as taught by Calle et al. in “Temporal analysis of tissue displacement induced by a transient ultrasound radiation force” (J Acoust Soc Am. 2005 November; 118(5):2829-40).
Steps 100 to 110 represent preparatory steps for the subsequent steps of the method of the invention, where the viscoelastic model will be used, as explained here-after.
Determination of Local Field Intensity
When the local intensity of the acoustic field AF is to be determined, one implements a further step 112 of the method where the acoustic field AF is applied to the target region 82.
Thereafter, the displacement u induced by the acoustic field is determined in a further step 114.
In step 114, a second displacement image DI′ of the soft solid in target region 82 is created. This second displacement image DI′, which is an image of the displacement u induced by the acoustic field, is determined by an ultrasound technique, as taught by Ophir et al in “Elastography: a quantitative method for imaging the elasticity of biological tissues” (Ultrason Imaging. 1991 April; 13(2):111-34) whose content is incorporated herein by reference.
With this ultrasound technique, estimation of the displacement induced by the ultrasonic beam to be measured involves the following elementary steps
Each position of volume V4 is observed in a coordinate system (x1,x2,x3).
Each pixel of an image I1 or I2 represents the local displacement undergone by this pixel along a chosen direction; for example along direction x3 in the coordinate system (x1,x2,x3), the direction being parallel to arrow AB.
Equation (8) shows that the entire displacement distribution ui,ss({right arrow over (r)}) is proportional to peak intensity I0. As a consequence, the experimental value of peak intensity can be estimated in three steps:
In practice, steady-state localized deformation at step 112 can be approximated by driving the acoustic array or source 2 with a burst having constant power and long duration T:
In other words, the rectangle function used to drive the acoustic source 2 at step 112 equals 1 between t=−T/2 and t=+T/2 and 0 otherwise. The rectangle function is a long burst between t=−T/2 and t=+T/2. In practice, the value of T is chosen between 0.1 and 1000 milliseconds (ms), preferably between 1 and 100 ms.
Alternatively, other functions can be used to define the shape and/or duration of application of the ultrasound beam. In particular, excitation can be made by a succession of short bursts, with a duration between 50 and 200 microseconds (μs) each, where the total duration of the bursts train is between 0.1 and 1000 ms, preferably between 1 and 100 ms.
The steady-state displacement is approximated by a measurement performed at step 114, immediately after the end of the excitation burst:
ui,ss({right arrow over (r1)})≈ui({right arrow over (r1)},T/2) (11)
By “immediately”, one means that measure of displacement takes place within 0 and 10 ms, preferably within 0 and 0.1 ms after the end of the application of the acoustic field.
Alternatively, the second displacement image DI′ is created at step 114 by MRI technique (Magnetic Resonance Imaging), as taught by Muthupillai et al in “Magnetic resonance elastography by direct visualization of propagating acoustic strain waves” (Science. 1995 Sep. 29; 269(5232):1854-7), as taught by McDannold and Meier “Magnetic resonance acoustic radiation force imaging” (Med Phys. 2008; 35(8):3748-58) or as taught by Souchon et al in “Transient MR elastography (t-MRE) using ultrasound radiation force: theory, safety, and initial experiments in vitro” (Magn Reson Med. 2008 October; 60(4):871-81), whose content is incorporated by reference.
According to another alternative approach, at step 114, the displacement can be measured by one or several optical methods.
Then, in a subsequent step 116, the intensity is computed on the basis of model M. Consider that a value um has been measured at a given position for displacement u at step 114. It is then possible to access model M in order to determine which value im of the intensity corresponds to value um, as shown by the dotted lines in
The goal of the invention is reached and one can, on the basis of the value computed in step 116, react to an attenuation or deflection of the beam B within target region 82.
Second Method of the Invention
The first method described here above, which is according to a first embodiment of the invention, gives good results.
However, it makes the assumption that the medium of the soft solid 8 is homogeneous. Such an assumption is not always valid, in particular for biological tissues.
In such a case, a possible solution consists in developing a model for heterogeneous media, for example using finite elements as taught by Palmeri et al in “A finite-element method model of soft tissue response to impulsive acoustic radiation force” (IEEE Trans Ultrason Ferroelectr Freq Control. 2005 October; 52(10):1699-712). However this approach requires exact knowledge of the viscoelastic properties of the tissues along the entire propagation path of the beam. In medical applications, these properties are usually unknown, or they are known with substantial uncertainty.
The second method of the invention represented on
In the description here below, steps of the methods and parts of the associated apparatus which are the same as in the first embodiment have the same references. Hereafter, only the differences with respect to the first embodiment are detailed.
Assume the force distribution {right arrow over (f)}({right arrow over (r)}) in any position {right arrow over (r)} of the target region can be separated in two parts, that is:
The following equation prevails:
{right arrow over (f)}({right arrow over (r)})={right arrow over (flocal)}({right arrow over (r)})+{right arrow over (fremote)}({right arrow over (r)}) (12)
A local region 82A corresponds to a small region of interest, where acoustic intensity is to be determined. This local region is a sub-part of target region 82. The local region 82A is assumed to be locally homogeneous. The remote region, that is the remaining part of soft solid 8 outside of local region 82A, can be heterogeneous. The steady-state displacement {right arrow over (uSS)}({right arrow over (r)}) can be decomposed as the sum of the displacement {right arrow over (uSS,local)}({right arrow over (r)}) induced by forces acting locally, that is in local region 82A, and of displacement {right arrow over (uSS,remote)}({right arrow over (r)}) induced by remote forces acting at a distance. The following equation prevails:
{right arrow over (uSS)}({right arrow over (r)})={right arrow over (uSS,local)}({right arrow over (r)})+{right arrow over (uSS,remote)}({right arrow over (r)}) (13)
Consider two observation positions {right arrow over (r1)} and {right arrow over (r2)} located within the local region of interest 82A, and close to each other. The steady-state displacement {right arrow over (uSS,remote)}({right arrow over (r2)}) induced by remote static forces at position {right arrow over (r2)} will be almost identical to the steady-state displacement {right arrow over (uSS,remote)}({right arrow over (r1)}) induced at position {right arrow over (r1)}. In other words, the following equation prevails:
Now, consider the difference {right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)}) defined by:
{right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)})={right arrow over (uSS)}({right arrow over (r1)})−{right arrow over (uSS)}({right arrow over (r2)}) (15)
Using the decomposition in local and remote contributions, as defined in equation 13, and noting that remote contributions are nearly identical, as shown in equation 14, equation (15) becomes:
{right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)})≈{right arrow over (uSS,local)}({right arrow over (r1)})−{right arrow over (uSS,local)}({right arrow over (r2)}) (16)
In conclusion, the difference in displacement {right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)}) undergone at two neighboring positions depends mostly on local forces. It is almost independent of the forces acting on the remote region.
The difference {right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)}) can therefore be calculated in a simulation wherein forces acting at a distance, in the heterogeneous remote region, are ignored. Only local forces acting within the homogeneous region of interest 82A need be accounted for. In this region, the homogeneous Green's function model of equation 8 can be used, and the difference in displacement {right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)}) can be expressed in the following equation:
Equation (17) shows that the difference {right arrow over (ΔuSS)}({right arrow over (r1)},{right arrow over (r2)}) is proportional to peak intensity I0.
The method of
In this second method of the invention, the experimental value of peak intensity can be estimated in four main steps:
Alternatively, in step 116, in the specific case of a linear model, as the one illustrated in
In practice, according to an advantageous aspect of the invention, in order to minimize the value of δ{right arrow over (uSS,remote)}, the positions {right arrow over (r1)} and {right arrow over (r2)} for which the displacement difference is considered are chosen at the same depth or approximatively at the same depth, with respect to the ultrasound source 2, within the target region 82. Consider the center C2 of ultrasound array 2, that is the center of the emission of beam B. One can choose that positions {right arrow over (r1)} and {right arrow over (r2)} are located at the same distance d of this center C2, that is, in the plane of
Preferably, one of the two positions {right arrow over (r1)} and {right arrow over (r2)} is chosen equal to the location {right arrow over (r0)} where the peak I0 of the intensity is expected to occur.
Optional Temperature Correction
For the two methods mentioned here above, local heating induced by the acoustic beam may result in artifacts in the displacement image obtained at steps 108 and 114. For small changes in temperature, typically within +/−10° C., this phenomenon is mostly due to speed of sound changing with temperature, as taught by Seip et al in “Noninvasive estimation of tissue temperature response to heating fields using diagnostic ultrasound” (IEEE Trans Biomed Eng. 1995 August; 42(8):828-39) or by Souchon et al in “Monitoring the formation of thermal lesions with heat-induced echo-strain imaging: a feasibility study” (Ultrasound Med Biol. 2005 February; 31(2):251-9), whose content is incorporated by reference.
Artifacts due to heating are corrected by an additional step 120 taking place between steps 114 and 116 of the second method, where the displacement due to heating is removed from the steady-state displacement value measured at step 114. More precisely, this step 120 includes the following elementary steps:
In other words, a first image I1 of the target region is obtained by ultrasound scanning prior to step 112. Then, one implements steps 112 and 114. Then one waits for the target region to return to its initial position which existed prior to step 112, as mentioned in elementary step viii). Thereafter, a third image of the target region is obtained by ultrasound scanning at the end of elementary step viii). A thermal displacement at the given position ({right arrow over (r)}) can then be determined by comparing the first and third images I1 and I3 respectively obtained. This corresponds to elementary step ix). Then, at step x), the steady-state displacement (ui,ss({right arrow over (r)})) measured at step 114 at a given position ({right arrow over (r)}) is corrected by subtraction of the value of the thermal displacement at the given position.
Alternatively, the corrected displacement is obtained directly by comparing the second image of the medium or soft solid 8, which was acquired during step 114, and the third image of the medium.
The time delay between step 114 and elementary step ix), that is the duration of step viii), must be long enough so that the medium or soft solid 8 has time to return to its initial position, and short enough so that cooling is negligible. A typical value for this time delay is between 1 and 100 ms.
Optional Phase Aberration Correction
The geometric distortion possibly undergone by the acoustic beam B, while propagating in the soft solid 8, is called phase aberration. This might induce some perturbation of the measured displacement.
Phase aberration can be corrected by applying time delays to the elements of the acoustic source. A set of time delays is called a delay law. The time delays are adapted to the medium, in other words to the soft solid 8. Such correction requires that the source is composed of multiple elements.
Thus, in a further step 122, phase aberration is detected based on the difference between the experimental displacement image obtained at step 114 and the displacement image expected from simulations for the set value of the spatial distribution of the acoustic field intensity computed with the model M′ at step 110. In this step 122, phase aberration is corrected iteratively, until the mismatch is minimized. Here, one can make use of the teachings of Herbert et al in “Energy-based adaptive focusing of waves: application to noninvasive aberration correction of ultrasonic wavefields” (IEEE Trans Ultrason Ferroelectr Freq Control. 2009 November; 56(11):2388-99) whose content is incorporated by reference. Iteration is shown by arrow 124 on
In other words, a measured displacement image can be generated for different given positions. Then, a possible phase aberration is detected, based on a mismatch between the measured displacement image and a displacement image expected from simulation. If is detected, phase aberration is corrected iteratively, until the mismatch is minimized.
On
General Remarks
The methods and apparatus described here above can be combined to generate new embodiments of the invention, within the scope of the attached claims. In particular, steps 120 and/or 122 can also be implemented with the method of
The invention is described here above in case one uses a Green's function to build the viscoelastic model M. However, other approaches can be performed to do so, such as the finite differences method, or using the finite elements method described by Eskandari et al. in “Viscoelastic characterization of soft tissue from dynamic finite element models” (Physics in medicine and biology 2008; 53(22): 6569-90).
Similarly, inversion of the viscoelastic model M can be performed with techniques alternative to the ones leading to equation (6). In particular, one can use any of the methods described in the book “Parameter Estimation and Inverse Problems”, (2018), Third Edition, Elsevier. ISBN 9780128134238, ISBN 9780128134238″ by Aster Richard; Borchers Brian, and Thurber Clifford, whose content is included in this description by reference.
As explained here above, a therapeutic ultrasound process can be implemented, wherein an ultrasound field is steered towards a target region of a biological tissue and wherein the local intensity of the acoustic field is determined with a method as explained here-above.
However, the invention is not limited to diagnostic or therapeutic ultrasound applications. It can be used to characterize different types of animal or vegetal soft solids and natural or synthetic soft material in cosmetic or food industry, in particular for quality control purposes. The invention can also be used to characterize non-metallic portions of prostheses. The invention can also be used to treat such soft solids, in particular to improve some of their physical characteristics.
Number | Date | Country | Kind |
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18305291 | Mar 2018 | EP | regional |
Filing Document | Filing Date | Country | Kind |
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PCT/EP2019/056549 | 3/15/2019 | WO |
Publishing Document | Publishing Date | Country | Kind |
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WO2019/175387 | 9/19/2019 | WO | A |
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20120226158 | Greenleaf | Sep 2012 | A1 |
Number | Date | Country |
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101785683 | Jul 2010 | CN |
1809377 | Apr 2017 | EP |
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---|
Bercoff, Jérémy, et al. “The role of viscosity in the impulse diffraction field of elastic waves induced by the acoustic radiation force.” IEEE transactions on ultrasonics, ferroelectrics, and frequency control 51.11 (2004): 1523-1536. (Year: 2004). |
Vyas, Urvi, Elena Kaye, and Kim Butts Pauly. “Transcranial phase aberration correction using beam simulations and MR-ARFI.” Medical physics 41.3 (2014): 032901. (Year: 2014). |
Green's function. Wikipedia. Wikimedia foundation. Accessed Jul. 8, 2022. Published Sep. 8, 2017. (Year: 2017). |
Palmeri Mike L et al “Evaluating the feasibility of acoustic radiation force impulse shear wave elasticity imaging of the uterine cervix with an intracavity array: a simulation study”, IEEE Transactions on Ultrasonics, Ferroelectrics and Frequency Control, IEEE, US, vol. 60, No. 10, Oct. 1, 2013. |
Number | Date | Country | |
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20210007713 A1 | Jan 2021 | US |