EP4007914A1 - Procédé et système de caractérisation ultrasonore non invasive d'un milieu hétérogène - Google Patents
Procédé et système de caractérisation ultrasonore non invasive d'un milieu hétérogèneInfo
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- EP4007914A1 EP4007914A1 EP20820244.0A EP20820244A EP4007914A1 EP 4007914 A1 EP4007914 A1 EP 4007914A1 EP 20820244 A EP20820244 A EP 20820244A EP 4007914 A1 EP4007914 A1 EP 4007914A1
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- A61B8/44—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device
- A61B8/4483—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device characterised by features of the ultrasound transducer
- A61B8/4488—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device characterised by features of the ultrasound transducer the transducer being a phased array
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- A61B8/5207—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of raw data to produce diagnostic data, e.g. for generating an image
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Definitions
- the present description relates to methods and systems for non-invasive ultrasound characterization of a heterogeneous medium, and applies in particular to medical imaging or to non-destructive testing and more generally to all the fields in which ultrasound imaging can be used. .
- the resolution of an acoustic imaging system can be defined as the ability to discern the small details of an object.
- an acoustic imaging system is limited by diffraction and the theoretical resolution is given by l / 2 (where l is the wavelength of sound in the medium), or by the finite angular aperture of the detector.
- the resolution is often deteriorated by variations in the speed of sound in the propagation medium.
- the medium is considered to be homogeneous, with a constant speed of sound cO.
- the hypothesis of a homogeneous environment is not always respected.
- the probe is placed between the patient's ribs. Acoustic waves travel through layers of fat and muscle before reaching the target organ. Soft tissues each have different mechanical properties. The speed of sound is therefore far from homogeneous, between 1450 m / s for adipose tissue and 1600 m / s for the liver.
- the variations in speed of sound cause a different phase shift of the waves depending on the places through which they propagate. This results in an aberration of the acoustic wavefront which leads to a distortion of the resulting ultrasound image, and therefore to a degradation of its resolution and its contrast. These aberrations can be such as to compromise the results of the medical examination.
- the first way to generate an echographic image of the medium to be studied is to emit an ultrasonic pulse from one of the transducers of the array, the position of which is identified by the vector u in (FIG. 1A, diagram on the left). This results in a diverging cylindrical (or spherical) incident wave for a 1D (or 2D) array of transducers. This is reflected by the diffusers 21 of the medium 20 and the backscattered field is recorded by each of the transducers 11 as a function of time (FIG. 1A, diagram on the right).
- FIG. 1C illustrates the principle of this so-called plane wave ultrasound, described for example in the article by G. Montaldo et al. “Coherent plane-wave compounding for very high frame rate ultrasonography and transient elastography” (IEEE Trans. Ultrason., Ferroelect. Freq. Control 56 489-506, 2009). Delays are applied to each signal on transmission (figure IC, diagram on the left) for the formation of a wavefront inclined at an angle q rapporth with respect to the array of transducers 10.
- the field backscattered by the medium, R (u ollt , 0 in , t) is measured by all the position sensors u out for a series of incident plane waves of which the angle of incidence q ⁇ h . All of these responses form a reflection matrix R ue (t) defined between the Fourier base (or plane wave base) at the input and the base of the transducers at the output. Once this matrix has been recorded, the signals are temporally shifted before being summed in a coherent manner in order to digitally focus the data on transmission and on reception for each point of position r in . The number of acquisitions necessary to form an echographic image is reduced compared to standard echography (focused emissions), and this for the same level of contrast and resolution of the echographic image.
- Figure 2 illustrates the influence of medium aberrations on conventional ultrasound imaging methods ( Figures IA to IC).
- the delays initially determined and to be applied to each of the transducers of the network on transmission and on reception are not optimal for the evaluation of an image of the medium, since they are determined with the assumption of a homogeneous medium with constant speed of sound.
- An aberration layer 22 induces a distortion of the incident wave front.
- step 25 the delay laws used do not make it possible to concentrate the acoustic energy in the zone delimited by the limits of the diffraction, zones usually called the focal spot.
- step 26 the delay laws used do not make it possible to correctly select the ultrasonic signals originating from the focal point of the medium, and mix the signals originating from a focal spot also aberrated. This results in a double aberration in the image construction process, which greatly degrades its resolution. New delay laws can then be recalculated in order to compensate for the effect of the aberration layer by adding an additional delay law to the delays generally used in channel formation.
- Patent application WO-2010/001027 proposes an ultrasonic probing method capable of separating the multiple scattering component from the simple scattering component by filtering a frequency transfer matrix representative of the responses between the transducers of all of the transducers. This method makes it possible to obtain information on multiple scattering in which the reflected background results from several successive reflections on diffusers, and for which the time of flight is not directly related to the distance between a diffuser and the transducers.
- the subject of the present description is a method of non-invasive ultrasonic characterization of a heterogeneous medium, the method comprising:
- the method makes it possible, in a very advantageous manner, to probe the medium very locally in any direction of the measurement axis passing through the first point and the second point, in order to determine, by the input and output focalizations, a new matrix of local responses REP (r, Ar) (very rich in local information) at any point of the midpoint of position r and for any angular direction b of analysis.
- the recorded experimental reflection matrix R ui (t) can be a “real” matrix, that is to say composed of real coefficients in the time domain, the electrical signals recorded by each of the transducers being real numbers.
- this matrix can be a “complex” matrix, that is to say made up of complex values, for example in the case of a demodulation for a formation of in-phase and in quadrature channels (known in English as the name “beamforming IQ”).
- the focusing at the input of the experimental reflection matrix uses a time of flight on the outward journey of the waves between the emission base and the transducer virtual input
- the output focusing uses a time of flight on the return of the waves between the virtual output transducer and the reception base.
- the response REP (r, Ar) of the medium is calculated by the following formula: in which
- Ni n is the number of elements of the emission base i
- N out is the number of elements of the reception base u at the output
- R Riveri (t) is the experimental reflection matrix, of which R Uj ( u out - ⁇ in J T ( r i n r out- u out- ⁇ ii l )) is the element of the experimental reflection matrix R lovedI (t) recorded by the transducer u out following the emission i in at time t, t is a time which is the sum of the time of flight on the outward journey t ih of the ultrasonic wave between the base of emission i and the first point PI and the return flight time x out of the ultrasonic wave between the second point P2 and the reception base u, as explained by the following formula:
- the method further comprises:
- a step of determining a profile of responses PR (ôr) which is a plurality of responses REP (r, Ar) calculated for a plurality of values of the distance coordinate Ar and for the same central point (PC), and for the same measurement axis (AX m ), corresponding to a predetermined angle b, ôr being the distance from the second point relative to the central point, that is to say the value such that Ar ôr.up, up being a unit vector in the direction of the measurement axis AX m defined by the angle b.
- the method further comprises:
- 0).
- the width of the peak is estimated at a height which is a portion of the maximum height of said peak, said portion being for example half of the maximum height of the peak.
- the method further comprises:
- the method further comprises:
- an image optimization step in which the focusing criterion F (r) is calculated at a plurality of points in the middle, and at least one parameter for calculating the input focusing and / or the output focusing is optimized. minimizing or maximizing an average of said focusing criterion F (r) for said plurality of points.
- said at least one calculation parameter comprises the speed of sound in the medium.
- the theoretical resolution is determined by a technique included in the following list:
- the method further comprises:
- the symmetry rate is calculated by the following formula: or by the following formula: in which
- the method further comprises:
- a noise intensity I ⁇ (r) being the product between one minus the rate of symmetry a (r) and the afocal intensity I, n ⁇ r), that is to say:
- the first multiple scattering indicator (r) is calculated by a ratio between the multiple scattering intensity I ⁇ i (r) and the noise intensity I ⁇ (r), that is - say:
- the method further comprises:
- a step of determining a confocal intensity I on (r) which is the value of a squared modulus of the response REP (r,
- 0) for a zero distance coordinate value (
- 0), i.e. for a point in the middle for which the first point (PI), the second point (P2) and the central point (PC) are the same,
- I on (r) IsO) + 2I M (r) + I N (r)
- the method further comprises: - a step of determining a second multiple diffusion indicator y (r) which is calculated by:
- the method further comprises:
- the local characterization parameter is chosen from a list comprising the resolution w (r), the focusing criterion F (r), the rate of symmetry a (r), the first multiple diffusion indicator (r ), the second multiple scattering indicator y (r), the afocal intensity I 0ff (i), the confocal intensity I on (r), the multiple scattering intensity IM (G), the scattering intensity simple Is (r), the noise intensity I N (r).
- the present description relates, according to a second aspect, to a non-invasive ultrasonic characterization system of a heterogeneous medium configured for the implementation of all the examples of ultrasonic characterization methods as described above.
- the ultrasonic characterization system according to the second aspect comprises:
- a computing unit (42) coupled to the first network of transducers and suitable for:
- the characterization system according to the present description can comprise at least one array of transducers which are both transmitter and receiver, or several arrays of transducers, some being dedicated to the transmission, others to the reception of ultrasonic waves.
- Figures IA to IC illustrate known transmission / reception mechanisms for ultrasound imaging and quantification
- Figure 2 illustrates the impact of aberrations in ultrasound imaging, according to the prior art
- FIG. 3 illustrates an example of an ultrasonic characterization system for the implementation of the ultrasonic characterization methods according to the present description
- FIG. 4 illustrates the definitions used in the ultrasonic characterization method according to the present description
- FIG. 5 illustrates an example of an image representing the modulus of a response matrix REP (r, Ar) according to the ultrasonic characterization method as in FIG. 4;
- FIG. 6 illustrates an example of a response profile PR (ôr) corresponding to the response matrix of FIG. 5;
- FIG. 7 illustrates echographic images of three heterogeneous media, part A of this figure corresponding to a test medium called "phantom", part B of this figure corresponding to a medium having a layer of meat placed on the medium of test, and part C of this figure corresponding to a medium which is an “in vivo” liver;
- FIG. 8 illustrates an image of resolution w (r) (on the left) and an image of theoretical resolution wo (r) (on the right) established for the test medium (of part A) of FIG. 7;
- FIG. 9 illustrates images of the focusing criterion F (r) established for the three heterogeneous media of FIG. 7 (respectively parts A, B, and C);
- Figure 10 illustrates images of the first multiple scattering indicator (r) of focusing established for the three heterogeneous media in FIG. 7 (respectively parts A, B, and C);
- FIG. 11 illustrates images of the second multiple scattering indicator y (r) established for the three heterogeneous media of FIG. 7 (respectively parts A, B, and C);
- FIG. 12 illustrates three calculations of optimum sound speeds, carried out from three models of the medium B of FIG. 7 and from the defined focusing criterion F (r).
- FIG. 3 illustrates an example of an ultrasonic characterization system 40 for the implementation of ultrasonic characterization methods of a heterogeneous medium 20, according to the present description.
- the system 40 comprises at least a first array 10 of transducers 11, for example a linear or two-dimensional array; the transducers are for example piezoelectric ultrasonic transducers which may conventionally be in the form of a rigid bar brought into contact with the medium 20.
- the array of transducers is for example part of a probing device 41, also referred to hereinafter as by the more common term "probe"; the array of transducers is connected to a computing unit 42, which can itself be connected to a display device 43; the computing unit transmits and records electrical signals to / from each of the transducers 11.
- connection or “link” between the probing device 41, the computing unit 42 and the display device 43, is meant any type of wire link of the electrical or optical type, or of wireless link using any protocol. such as WiFi, bluetooth or others. These connections or links are one way or two way.
- the calculation unit 42 is configured for the implementation of calculation or processing steps in particular for the implementation of process steps according to the present description.
- a spatial reference of the medium 20 by taking a first X axis and a second Z axis perpendicular to it.
- the first X axis corresponds to the direction in which the transducers 11 are aligned for a linear array
- the second Z axis corresponds to the depth of the medium 20 with respect to this array 10 of transducers 11.
- This definition can be extended. to a three-axis spatial coordinate system in the case of a two-dimensional network.
- FIG. 3 as in the rest of the description, reference is made to an array of transducers for transmission and reception, it being understood that, in a more general case, several arrays of transducers could be used simultaneously. They can be both sender and receiver, or only a sender for some and only a receiver for others.
- each calculation or processing step can be implemented by software, hardware, firmware. , microcode or any suitable combination of these technologies.
- each computational or processing step can be performed by computer program instructions or software code. These instructions can be stored or transmitted to a storage medium readable by a computer (or calculation unit) and / or be executed by a computer (or calculation unit) in order to implement these calculation or processing steps.
- a central point PC of spatial position r in the spatial reference frame of the medium is defined, and situated in the middle of a first point PI and of a second point P2.
- a measurement axis AX m passing through the first point PI and the second point P2, and forming an angle b with respect to the first axis X of the array of transducers 11.
- the central point PC is located on the origin of the.
- measurement axis AX m zero distance coordinate on the measurement axis).
- the first point PI is at a distance coordinate -Ar and the second point P2 is at a distance coordinate + Ar from the central point PC, origin of the measurement axis.
- the spatial position r and the distance coordinate Ar are noted in bold, signifying that these elements are position and offset vectors with respect to a position, vectors taken from the spatial reference frame of the medium (X, Z).
- the coordinate vector distance DG thus takes into account the direction of the measurement axis AX m , and its angle b with respect to the first axis X.
- Other definitions of the positions of the points relative to the others are possible and accessible to any specialist in the trade ultrasound.
- the first and second points can be identified by a distance
- These two points PI and P2 can be a short distance from each other, that is to say a few millimeters from each other, and for example 20 millimeters or less.
- the method of non-invasive ultrasound characterization comprises:
- the emission base i at the input being for example a base of waves each generated by only one of the transducers 11 of the array 10 or a base of plane waves of angular inclination Q relative to the axis X, as described previously in the description of Figures IA to IC.
- the reception base u is usually the base of the transducers 11.
- the step of generating ultrasonic waves is understood between the transmission base i and the reception base u.
- This ultrasonic generation step is therefore defined for any type of ultrasonic wave of the focused or unfocused type, such as plane waves.
- the experimental reflection matrix R beaui (t) is therefore defined between the transmission base i at the input and a reception base u at the output.
- This matrix contains all the temporal responses of the medium, measured at time t by each transducer 11 with spatial coordinate u out for each emission i in . It is understood that the elements named with the index "in” refer to the transmission (ie the entry) and the elements named with the index "out” refer to the reception (ie the exit).
- the input focusing process uses a time of flight on the outward journey of the waves between the emission base (i) and the virtual input transducer TV in .
- the focus-out process uses a return time-of-flight of the waves between the virtual output transducer TV out and the transducers at the receiving base (u).
- the first point PI being relative to the virtual input transducer TVi n , it is then located at a coordinate -Ar on the measurement axis AX m with respect to the central point PC and the second point P2 being relative to the virtual output transducer TV out , it is then located at a + Ar coordinate on the measurement axis AX m with respect to the central point PC.
- the input focusing (on emission) is configured to concentrate the acoustic wave around the point PI on a spatial extent corresponding to the input focal spot.
- the diffusers located inside this zone in the medium then generate a wave which is backscattered towards the probe.
- This zone characterized by the focal spot on emission and the reflectivity of the corresponding medium, is called the virtual input transducer TVi n and can be interpreted as a "virtual" source.
- the output focusing (on reception) is configured to select the echoes generated by diffusers located around the point P2 over a spatial extent corresponding to the output focal spot.
- This zone characterized by the output focal spot (reception) and the reflectivity of the corresponding medium is called the virtual output transducer TV out and can be interpreted as a "virtual sensor”.
- the response REP (r, Ar) can therefore be interpreted as an estimate of the pressure field coming from the position r out for a focusing at the position r in .
- the virtual input transducer TVi n corresponds to an ultrasonic "virtual source" at the spatial position r; "in the medium and the virtual output transducer TV out corresponds to a" virtual ultrasonic sensor at the spatial position r out .
- the method is suitable for probing the medium around point PI and / or point P2, spatially, which makes it possible to obtain new information on the propagation of the waves.
- Ni n is the number of elements of the emission base i
- N out is the number of elements of the reception base u at the output
- R Riveri (t) is the experimental reflection matrix, of which R Uj ( u out - ⁇ i n J T ( r i n r out - u out - ⁇ i l )) is the element of the experimental reflection matrix R stunti (t) recorded by the transducer u out following the emission i; parcel at time t.
- the time t is the sum of the time of flight on the outward journey t ih of the ultrasonic wave between the transducers of the emission base i and the first point PI and the time of flight at the return x out of the ultrasonic wave between the second point P2 and the transducers of the reception base u, as explained by the following formula:
- the flight times t ih and x out are calculated from a speed of sound model.
- the hypothesis consists in considering a homogeneous medium with a constant speed of sound ⁇ 3 ⁇ 4. In this case, the flight times are obtained directly from the distances between the probe and the virtual transducers.
- the number of elements of the emission base Ni n is for example greater than or equal to two.
- the number of elements of the reception base N out is for example greater than or equal to two.
- This improved path formation formula is therefore a double sum of the temporal responses recorded in the experimental reflection matrix R réellej, a first sum related to the input focusing (the emission) according to the emission base i at the point PI of spatial position ,, and a second sum related to the output focusing (reception) according to the reception base u at point P2 of spatial position r out.
- This calculation is therefore carried out for the spatial coordinates of the two points PI and P2 (of spatial positions r in , r out ).
- the result of this improved path forming formula is therefore a temporal signal which is a pressure field for these two spatial coordinates (h handed, r out ).
- the method makes it possible to probe the medium very locally in any direction corresponding to the measurement axis AX m , in order to extract from it by the input and output focalizations more local information on the medium at the central point PC of spatial position r, between the first point and second point of the heterogeneous medium 20.
- This local information is entirely contained in the values of the calculated response, the response REP (r, Ar) of the medium which can be used to characterize each midpoint, for example in terms of resolution or in terms of multiple scattering.
- This local information is entirely contained in the values of the calculated temporal response and which can be used to characterize each point in the middle.
- the responses REP (r, Ar) can thus be determined for any set of separation distance values
- , b) are then the polar coordinates of the vector distance coordinates Ar.
- the response REP (r, -Ar) corresponds to inverting the spatial positions of the input and output virtual transducers.
- the set of responses REP (r, Ar) can then be recorded in a matrix of the same name. This matrix of responses in a focused reflection matrix, which records a pressure field calculated at any point in the middle with the defined hypotheses.
- FIG. 5 shows an image corresponding to a sub-matrix of the response matrix REP, said sub-matrix corresponding to a set of several central points PC of spatial position r in which the Z axis coordinate is fixed, and the angle b is zero.
- the abscissa corresponds to a variation along the X axis of the position of the central point PC and the ordinate corresponds to the distance Ar coordinate from this central point.
- the values of the points of this image (response) outside the x-axis of this image have a low (but not zero) value.
- the values of the points of this image (response) on the abscissa axis have a value corresponding to the intensity of the ultrasonic image point at the central point PC.
- Such an image can be extracted from the response REP (r, Ar) for variations in the distance coordinates Ar on a single measurement axis AXm or several measurement axes, that is to say for one or more values d 'angle b.
- and the angle b can also be constructed, which gives a representation of the variation of the response around a central point PC, and therefore of the focal task at that point.
- a step of determining a response profile PR (ôr) can be carried out, the response profile being a plurality of the responses REP (r, Ar) calculated for a plurality of values of the distance coordinate Ar.
- This response profile PR (ôr) is considered for the same central point PC of spatial position r and along the same measurement axis AX m , corresponding to the same direction of angle b .
- the profile of responses PR (ôr) is a vertical section of the image of FIG. 5, and this response profile is the curve of said image section.
- the REP (r, Ar) responses can be complex values in particular in the case of use of formulation of focusing in complex values, as that is known in formations of channels in phase and in quadrature (known in English under the name "Beamforming IQ"). Consequently, the response profile PR (ôr) can also be represented by any module of these complex responses.
- PR response profile
- angle b can take any value between zero (0) and pi (p), and therefore a response profile PR C curve can be plotted or determined for several values of angle b.
- the set of response profiles PR (ôr) or PR (ôr, b) (if we take several angles, but we will only keep the spatial position in the following for the sake of simplification of the presentation) or PR ( r, ôr, b) (if we also take the spatial position of the central point) can be recorded in a matrix of the same name.
- the PR response profile (ôr) presents:
- the sub-matrix represented in FIG. 5 is therefore the set of response profiles PR (ôr) for a set of central point PC of spatial position r in which the axis coordinate Z is fixed, and the angle b is zero (or constant).
- This PR response profile (ôr) is a basic representation making it possible to determine new parameters for local characterization of the medium and / or of the performance of the ultrasound imaging process (ie of the formation of a pathway).
- ôr This PR response profile
- the modulus of the response profile as represented in FIG. 6 comprises a peak or maximum around the zero distance Ar coordinate (
- 0).
- This peak or over-intensity of the response profile is linked to the single scattering echoes and appears when the focal spots of the two virtual transducers TVi n and TV out overlap.
- the spatial extent of this peak is therefore strongly correlated with the spatial dimensions of the focal spots at the entrance and at the exit along the direction of angle b and therefore with the local resolution of the echographic image.
- the resolution w (r) can then be determined for example by the width of this peak.
- the width of this peak is for example determined at a height which is a portion of the maximum height of this peak. For example, that portion of the pitch will be half or a third (1/3) or two-thirds (2/3) or any other ratio of the maximum pitch.
- the maximum height of the peak is in fact the intensity of the echographic image at the central point PC if we look only at the squared modulus of the response profile, ie
- PR (ôr 0)
- the method makes it possible to define at any point in the middle the extent of the focal spot and therefore the resolution of the ultrasound process in each of the angle b directions.
- the image in the left part of Figure 8 shows an example of calculating the resolution w (r, b) at each point of the test midpoint (midpoint A). Note that the resolution degrades with depth and when moving towards the edge of the image.
- the theoretical resolution wo (r) is determined by a first analytical calculation at the central point (PC) for a pulse (coi), the emission base (i) and the reception base (u): It is calculated by the angle from which the transducer array is viewed from the central point (PC). It depends on the maximum angular half-opening used in transmission to insonify the central point of spatial position r or in reception to collect the echoes coming from this central point.
- the theoretical resolution wo (r) is determined by a second analytical calculation at the central point (PC) for a pulse range (Dw) and the emission base (i) and the reception base (u) . It is obtained by an integral calculation over said pulse range and the angle from which the transducer array is viewed from the central point (PC) weighted by the frequency spectrum of recorded signals.
- the latter can be obtained by averaging the modulus of the Fourier transform of the elements of the experimental reflection matrix R rebooti (t).
- the theoretical resolution wo (r) is determined by a third wave propagation simulation calculation, firstly between the first point in the middle corresponding to a virtual input transducer (T Vi n ) and the base of emission (i), and secondly between the second point in the middle corresponding to a virtual output transducer (TV out ) and the reception base (u), said simulation using the response REP (r, Ar) and a propagation model of waves in the middle.
- This third calculation reflects the double focusing step performed in order to calculate the response profiles PR (r, b, ôr).
- This third simulation calculation consists in generating a theoretical reflection matrix associated with a random medium whose speed of sound corresponds exactly to the speed of sound model assumed to calculate the responses REP (r, Ar).
- This simulation then uses the same transmission base and the same reception base as those used for the physical experiment. All the operations performed to determine the resolution w (r) are then repeated to calculate the theoretical resolution wo (r) from the generated theoretical reflection matrix. All the diffraction phenomena are perfectly taken into account and an estimate of the theoretical resolution of the medium without aberration is thus obtained. Note that the statistical properties of the medium such as the mean reflectivity of an area, the spectrum of the retro-scattered echoes can be deduced from the responses REP (r, Ar) in order to use a simulation which best models the experience carried out .
- the image in the right part of Figure 8 shows an example of the calculation of the theoretical resolution wo (r) at each point of the test medium (medium A). We notice that the resolution degrades with depth and also when moving towards the edge of the image. Note that this image is very similar to the image on the left side of this same figure. The calculation of the resolution carried out by the preceding method is therefore in agreement with the calculation of the theoretical resolution.
- the theoretical resolution is for example determined from an emission base i at the input, a reception base u at the output and a modeling of the propagation of ultrasonic waves in the medium.
- the focusing criterion F (r) is a ratio of said resolution and theoretical resolution, or the reverse (a simple convention). That is, that we can get:
- FIG. 9 illustrates images of the focusing criterion F (r) established for the three heterogeneous media in FIG. 7 (media A, B, and C in correspondence between FIGS. 7 and 9).
- a value of one (1) for this focusing criterion corresponds to an identical resolution and a theoretical resolution (in the clear in the figure).
- a value close to zero (0) for this focusing criterion corresponds to divergent resolution values (in dark in the figure), that is to say a degraded focusing.
- the image of the medium A illustrates a great homogeneity with an average of this focusing criterion close to 0.97. This means that the ultrasound image is well formed and that the focusing assumptions are correct.
- the images of media B and C show notable degradations corresponding to heterogeneities located upstream of the propagation of ultrasonic waves: layer of meat on the surface for medium B and adipose or muscle tissue for the liver of medium C.
- the image of the middle C highlights very degraded areas (dark areas at the bottom left of this image) which means that the image produced in FIG. 7 (part C) also has the same problems in the same places.
- the image of the focusing criterion can indicate to the operator of the echographic system that the left part of his echographic image is of poor quality, in particular in spatial resolution. This can help him to interpret his echographic image to establish a diagnosis preferentially on the correctly imaged areas or to modify the way of generating said echographic image. For example, this can encourage the operator to modify the positioning of the probe (ie the array of transducers) so as to obtain a good quality focusing criterion in the area (s) of interest for medical diagnosis. .
- the imaging system which implements the present technique will be able to superimpose an echographic image on an image of the focusing criterion.
- the average correlation coefficient is averaged for modulus distance Ar coordinate values greater than a predetermined resolution wa (r) (as shown in Figure 6) and / or being performed for a range of values d 'angle b or for a predetermined angle value b 0
- the predetermined resolution wa (r) is advantageously a value greater than half of the resolution w (r).
- the predetermined resolution wa (r) is a value greater than once, twice or three times the resolution w (r), in order to better exclude the values of the peak, as can be seen in FIG. 6.
- the average of the correlation coefficient according to this variable means that the correlation coefficient is averaged for values of the distance coordinate Ar far from the peak, and therefore for virtual transducers far from the central point PC, which in fact makes it possible to test the acoustic reciprocity of this central point PC, ie the symmetry of the response matrix REP (r, Ar) with respect to the abscissa axis (see example of such a matrix in FIG. 5).
- values beyond the predetermined resolution are small, comparison by an average correlation of symmetry can reliably estimate a symmetry rate a (r) of the central point PC.
- the correlation coefficient is averaged for a range of values of angle values b or for a predetermined angle value b ⁇ i.
- the average of the correlation coefficient according to this variable means that we average the correlation coefficient for one or more angle values b, which makes it possible to test the angular symmetry or according to an angular sector of the symmetry rate around of the central point PC, ie the reciprocity of this central point PC.
- the correlation coefficient is averaged for values of distance coordinates Ar with a modulus greater than a predetermined resolution Wd (r). and for a range of angle values b.
- the correlation is averaged over distance coordinate values for which the simple diffusion contribution is zero, the latter only appearing for a modulus of the distance coordinate lower than the predetermined resolution. It also makes it possible to obtain a more robust and stable estimate of the symmetry rate a (r).
- ⁇ > is the mathematical average operator, this operator being able to be performed according to one or more variables (for example above distance coordinate values greater than a predetermined resolution and values of angle b), and * is the conjugation operator for complex numbers.
- the rate of symmetry a (r) close to zero (0) if the propagation of the ultrasonic waves does not behave reciprocally around the central point PC, and the rate of symmetry a (r) close to one (1) whether the ultrasonic wave propagation behaves symmetrically or reciprocally around the central point PC.
- this symmetry rate tests the validity of the acoustic reciprocity for the part of the signals corresponding to a multiple scattering. This makes it possible to discriminate the multiple scattering of the noise, the noise not respecting the property of acoustic reciprocity.
- This first multiple diffusion indicator (r) is zero if the symmetry rate is zero, and it tends to infinity if the symmetry rate is close to one (1).
- Figure 10 illustrates images of the first multiple scattering indicator (r) established for the three heterogeneous media of Figure 7 (media A, B, and C in correspondence between Figures 7 and 10). These figures are graduated in decibels (logarithmic scale).
- this first multiple scattering indicator shows a localized zone of diffusion of ultrasonic waves behind one of the cylinders of heterogeneity for medium A.
- this first multiple diffusion indicator highlights the diffusion in a majority the volume of these media, either downstream of the meat part of medium B, or downstream of the fatty tissue of the liver.
- I (r, DG) is the intensity of the point, ie the squared modulus of the response at this point,
- is the mathematical operator of modulus, and ⁇ .> is the mathematical average operator, this operator being able to be carried out according to one or more variables (for example above distance coordinate values greater than a predetermined resolution and values of angle b).
- the afocal intensity I 0ff (r) makes it possible to characterize the incoherent energy in the echographic image.
- the afocal intensity I 0ff ( r ) results from the contribution of echoes originating from multiple scattering phenomena (which are reciprocal) or originating from electronic noise (which are of random type). It is therefore possible to decompose the afocal intensity I 0ff (r) into a multiple scattering intensity and a noise intensity using the symmetry rate a (r).
- the first multiple scattering indicator (r) can also be calculated by a ratio between the multiple scattering intensity I M (r) and the noise intensity I ⁇ (r), i.e. according to the following formula:
- I on (r) I s (r) + 2I M (r) + l N (r)
- a step of determining a second multiple diffusion indicator y (r) can then be carried out, this second multiple diffusion indicator being calculated by the following formula:
- FIG 11 illustrates images of the second multiple scattering indicator y (r) established for the three heterogeneous media of Figure 7 (media A, B, and C in correspondence between Figures 7 and 10). These figures are graduated in decibels (logarithmic scale).
- this first multiple scattering indicator corresponds to a large proportion of multiple scattering compared to single scattering (in the clear in the figure).
- a low value for this second multiple scattering indicator corresponds to little multiple scattering compared to single scattering (dark in the figure).
- this second multiple scattering indicator shows that the medium C is very scattering.
- this second multiple scattering indicator makes it possible to determine a new criterion sensitive to the phenomenon of multiple scattering, independently of the attenuation of the environment.
- the echographic image is for example calculated by successive focalizations at the input and the output of all the points in the medium.
- the set of points of the axis abscissa of FIG. 5 corresponds to a line of this echographic image calculated at the depth z used to form the image of this FIG. 5.
- this ultrasound image calculation strongly depends on an assumption of a homogeneous medium in which the speed of sound (the speed of propagation of ultrasonic waves) is well known and constant. If this assumption is incorrect, the focusing delay laws do not correspond to the medium considered and the focusing is imperfect. The resolution of the ultrasound image is then degraded. Other focus calculation parameters may influence focus.
- An image optimization step can then be carried out in which the focusing criterion F (r) is calculated at a plurality of points in the middle (for example a predetermined area of the image), and at least one parameter is optimized. of calculating input focusing and / or output focusing by minimizing or maximizing an average of said focusing criterion F (r) for said plurality of points.
- said at least one calculation parameter comprises at least the speed of sound in the medium.
- this calculation parameter is the speed of sound.
- the quality of focusing and therefore the focusing criterion is maximum when the speed of sound model used to carry out the focusing coincides with the true speed of sound of the medium.
- each layer being assumed to be homogeneous with a speed of sound Ci of this layer. It is necessary to estimate the thickness of each layer, either with a priori knowledge of the environment, or using a first ultrasound image. We then estimate by optimization of the focusing criterion the speed of sound in the outermost layer with a homogeneous model, then we estimate by optimization the next layer the speed of sound in the next layer with a two-layer model, and and so on.
- FIG. 12 illustrates the optimizations carried out for the same predetermined zone, the medium B by using: a) a model with a curved layer B1 with an optimal speed of sound cl opt , b) a B2 curve two-layer model with optimal sound speed c2 opt , c) a B1 curve three-layer model with optimal sound speed c3 opt ,
- the optimization method presented above and the multilayer modeling of the medium therefore make it possible to refine the speed of sound estimates, and therefore make it possible to determine the evolution of the speed of sound in the depth of the medium.
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