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WO2002056666A2 - Procede de detection d'un agent contraste ultrasonore dans un tissu mou, et de quantification de la perfusion sanguine a travers des regions tissulaires - Google Patents

Procede de detection d'un agent contraste ultrasonore dans un tissu mou, et de quantification de la perfusion sanguine a travers des regions tissulaires Download PDF

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Publication number
WO2002056666A2
WO2002056666A2 PCT/NO2001/000022 NO0100022W WO02056666A2 WO 2002056666 A2 WO2002056666 A2 WO 2002056666A2 NO 0100022 W NO0100022 W NO 0100022W WO 02056666 A2 WO02056666 A2 WO 02056666A2
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WIPO (PCT)
Prior art keywords
tissue
transmit
contrast agent
acoustic
pulse
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PCT/NO2001/000022
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English (en)
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WO2002056666A3 (fr
Inventor
Bjørn A. J. ANGELSEN
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Priority claimed from US09/766,328 external-priority patent/US6461303B2/en
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Publication of WO2002056666A2 publication Critical patent/WO2002056666A2/fr
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Classifications

    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/48—Diagnostic techniques
    • A61B8/481—Diagnostic techniques involving the use of contrast agents, e.g. microbubbles introduced into the bloodstream
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01S—RADIO DIRECTION-FINDING; RADIO NAVIGATION; DETERMINING DISTANCE OR VELOCITY BY USE OF RADIO WAVES; LOCATING OR PRESENCE-DETECTING BY USE OF THE REFLECTION OR RERADIATION OF RADIO WAVES; ANALOGOUS ARRANGEMENTS USING OTHER WAVES
    • G01S7/00—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00
    • G01S7/52—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00 of systems according to group G01S15/00
    • G01S7/52017—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00 of systems according to group G01S15/00 particularly adapted to short-range imaging
    • G01S7/52023—Details of receivers
    • G01S7/52036—Details of receivers using analysis of echo signal for target characterisation
    • G01S7/52038—Details of receivers using analysis of echo signal for target characterisation involving non-linear properties of the propagation medium or of the reflective target
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01S—RADIO DIRECTION-FINDING; RADIO NAVIGATION; DETERMINING DISTANCE OR VELOCITY BY USE OF RADIO WAVES; LOCATING OR PRESENCE-DETECTING BY USE OF THE REFLECTION OR RERADIATION OF RADIO WAVES; ANALOGOUS ARRANGEMENTS USING OTHER WAVES
    • G01S7/00—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00
    • G01S7/52—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00 of systems according to group G01S15/00
    • G01S7/52017—Details of systems according to groups G01S13/00, G01S15/00, G01S17/00 of systems according to group G01S15/00 particularly adapted to short-range imaging
    • G01S7/52023—Details of receivers
    • G01S7/52036—Details of receivers using analysis of echo signal for target characterisation
    • G01S7/52038—Details of receivers using analysis of echo signal for target characterisation involving non-linear properties of the propagation medium or of the reflective target
    • G01S7/52039—Details of receivers using analysis of echo signal for target characterisation involving non-linear properties of the propagation medium or of the reflective target exploiting the non-linear response of a contrast enhancer, e.g. a contrast agent

Definitions

  • the present invention is directed to a method for detecting an ultrasound contrast agent in a soft tifs ⁇ ue and quantitating blood perfusion through regions of tissue by detecting the contrast agent in the tissue.
  • An ultrasound contrast agent is a solution of small gas bubbles (diameter ⁇ 5 ⁇ m) that is injected into the blood stream. Such bubbles show strong and nonlinear scattering of ultrasound at the frequencies used for medical ultrasound imaging. Medical applications of the contrast agents include, but are not. ] irn.i ted to, enhancing imaging of blood vessels, improving the 0 detection of the endocardium as a border of the ventricular cavities, and improving the detection of blood jets through leaking cardiac valves or septal def cts .
  • Second harmonic ultrasound imaging is today 5 the commonly used method for detecting and imaging • ultrasound contrast agent in the tissue.
  • the non-linear elastic properties of the contrast agent bubbles produce higher harmonic components and sub harmonic components of the transmitted pulse frequency band in the scattered signal directly in the scattering process.
  • wi h the present wideband transducer technology it is only possible to utilize the second harmonic component of the signal by transmitting an ultrasound pulse with frequency spectrum in the lower part of the active frequency band of a wideband transducer.
  • the second harmonic component of the scattered signal io then received in the upper part of the transducer frequency band.
  • the non-linear elasticity of the contrast agent is much stronger than that of the tissue. Accordingly, considerable distortion of the scattered pul se directly in the scattering process results with high amplitudes in the 3rd and 4th harmonic component of the incident pulse frequency band. More importantly, the scattered amplitudes from the contrast agent in these frequency bands are much stronger than the scattered amplitudes in the 3rd and 4th harmonic frequency bands from the tissue. Therefore, the use of harmonic frequency bands higher than the 2nd component of the transmitted pulse frequency band for detection and imaging of the contrast agent provides improved separation between the signal amplitudes from the contrast agent and the signal amplitudes from the tissue.
  • the first problem is that the present medical ultrasound transducers have so narrow a bandwidth that it is not possible to transmit a pulse with frequency band around f 0 , and receive back-scattered frequency components in the frequency bands around 0 and 4f u with adequate sensitivity.
  • Adequate wideband transducers have bee made by highly damping the transducers, but this reduces the sensitivity to the signal scattered from the contrast agent in the myocardium below tolerable limits.
  • a transducer is used with capabilities of transmitting frequencies in a band around f 0 , with high sensitivity in the receive band around the 3rd or 4th harmonic component of the transmit band.
  • the high receive sensitivity is obtained by using resonant operation of the transducer in the receive band with minimal dampening .
  • the transmitted pulse must. have sufficiently limited amplitude in the receive frequency bands.
  • a solution to this problem is presented according to the invention by either bandpass filtering the transmitted pulse both in the transducer and/or electrically before driving the transducer, or by using band limited pulse generator with linear drive, amplifiers of the array transducer elements .
  • a second problem associated with utilizing the 3rd and 4th harmonic component of the transmitted frequency band for detection and imaging of ultrasound contrast agent is that the pulse distortion in the f> scattering from the contrast agent bubbles, highly depends on the amplitude of the pulse incident on the bubble.
  • the power absorption in the tissue 0 considerably reduces with the frequency, being approximately 0.5 dB/cmMHz.
  • the total absorption attenuation from 2-10 cm is ⁇ 3.5 d ⁇ .
  • Geometric focussing of the beam to the far 5 end of the image range may be used to compensate for this absorption attenuation. Due to diffraction at such low frequencies, the maximal amplitude in the focussed beam is found closer to the transducer than the geometric focus.
  • the transmit beam focussing will give a gain of ⁇ 3,4 dB from 2 -10cm with a 18mm circular aperture.
  • a transmit center frequency of 0.875MHz gives 3rd and 4th harmonic center frequencies at 2.625MHz and 3.5MHz, which are typical frequencies used for cardiac imaging. These frequencies produce tolerable absorption attenuation of the backscattered signal so that it can be compensated for by a depth variable receiver gain.
  • a receive frequency in the range of 2.5 - 4MHz also gives a lateral resolution of the receive beam comparable to that with regular echocardiography .
  • the depth along the receive beam is divided into sub-ranges.
  • Each sub-range is observed at different time intervals with different transmit pulses, where the focus of each transmit pulse is located within the corresponding receive range, and both the transmit focus, the transmit amplitude, and the transmit aperture are adjusted for optimal equalization of the incident pulse amplitude within the corresponding receive range, under the actual absorption of the ultrasound in the tissue.
  • the range could be divided into subranges from 2-6 cm to be interrogated with a transmit pulse with focus at 7cm with a reduced transmit aperture to secure sufficient focal depth, followed by a 2nd transmit pulse focussed at say 18 cm, using larger transmit amplitude and transmit aperture to achieve the same incident pulse amplitude at 2 and at 15 cm.
  • Further detailed optimization of number of transmit pulses, with corresponding transmit foci, amplitudes, and apertures can be done within the scope of the invention. Pulse destruction by the transmit pulses must also be taken Into the account in this optimization, as described below.
  • a third prob em associated with utilizing the 3rd and 4th harmonic component of the transmitted frequency band for detection and imaging of ultrasound contrast agent is that the power of the received signal from a depth, is proportional to the concentration of contrast agent in the tissue.
  • the source of contrast agent in a region of the myocardium is the inflow of blood to the region with a sink of contrast agent produced by the venous outflow.
  • the concentration of contrast agent in a tissue region is hence given by the product of the blood concentration in the region and the concentration of contrast agent in the inflowing blood.
  • the contrast agent concentration will in this situation not be reduced before close to complete blocking of the blood flow to the region occurs. Therefore, in a stationary situation, the received signal power from a range will give limited quantitative grading of the perfusion through the tissue region.
  • One method to improve the grading of the perfusion is to introduce an additional sink of contrast agent in the region such, for example, as by destroying the contrast agent with high amplitude incident ultrasound pressure pulses. Partial destruction of the contrast agent will give a concentration that depends on both the inflow rate of blood to the region (i.e., the source of contrast agent) and the destruction rate (i.e., the additional sink of contrast agent) . The concentration of contrast agent will in this case quantitatively be reduced with reduced blood perfusion in the tissue. By comparing the signal level from one region with the signal level from regions with normal perfusion, the received signal level will give a quantitative measure of regionally reduced perfusion. With complete destruction of the contrast agent in the image region, one can use the re-filling time of contrast agent in the tissue as a quantitative measure of the perfusion through the tissue.
  • the amplitude of the destruction pulses can be depth tailored as for the image pulses, using multiple transmit pulses with optimized transmit foci, amplitudes, and apertures.
  • the pulse with focu3 in one sub-range will then provide some contrast agent destruction at other ranges, and the whole set of transmit pulses for each image direction, must be designed for equal and limited destruction of the contrast agent along the whole image depth.
  • Multiple foci destruction pulses will also provide minimal width of the destruction beam for all depths, ensuring that the destruction pulses for one image beam produce limited contrast agent destruction in neighboring image beams.
  • Destruction pulses with narrow focus at small depths are then usod first to destroy contrast agent at low ranges.
  • the rapid geometric widening of the beam past the focus wili then produce an attenuation of the destruction pulse which reduces contrast agent destruction at larger depths.
  • Using a higher frequency of the shallow range destruction pulses will also increase the attenuation of these pulses at deeper ranges.
  • the destruction of the contrast agent at deeper ranges is then followed up with new destruction pulses with deeper foci, larger transmit amplitudes and apertures, and possible lower center frequency.
  • the overlap of destruction from the pulses at all ranges must be taken into account. Therefore, the whole sen of destruction pulses must be designed for each receive beam direction so that even destruction of the contrast agent over the whole image range occurs .
  • Contrast agent destruction by neighboring beams will only complicate the imaging at the edges of the scan, where the edge beam has a single neighbor.
  • a method for detection of ultrasound contrast agent in soft tissue includes utilizing an ultrasound transmit beam former and transducer array assembly for transmitting directive, focussed ultrasound pressure pulses with steerable transmit amplitude, transmit aperture, transmit focus, and transmit direction, and with temporal frequency component ' s within a limited band B centered at frequency f , towards a region of soft tissue that contains ultrasound contrast agent bubbles.
  • the transmit pulse parameters are arranged, possibly using multiple transmit pulses, so that the incident pressure pu3 se that is utilized for imaging of the contrast agent for a particular depth, has minimal variation over the actual image range.
  • the non-linearly distorted, back- scattered ultrasound signal is received from both the tissue and the ultrasound contrast agent bubbles with the same ultrasound transducer assembly and the received array element signals are passed through a receiver beamfor er that has a stearable spatially directive receiver sensitivity.
  • the transducer assembly has high sensitivity at the receive band of frequencies centered at 3f ⁇ and/or 4f Q for maximal sensitivity of the distorted, non- linearly scattered signal from the contrast agent bubbles.
  • the received signal is high-pass filtered so that the lowest frequency component of the resulting signal is at east 2 times higher than the frequency component of the transmitted signal.
  • the amplitude of the high-pass filtered signal is used for detecting ultrasound contrast agent bubbles buried within the tissue, and for imaging of contrast agent bubbles in the tissue.
  • the depth variation of the amplitude of the incident pressure pulse is minimized by positioning the transmi t focus deeper than the image range .
  • the width of the incident beam at each location is reduced, and the depth variation of and amplitude of the incident pressure pulse is minimized by dividing the total imaged depth range into sub-ranges, where a separate transmit pulse is used to interrogate each sub-range consecutively in time, arranging the transmit focus, the transmit aperture, and the. transmit amplitude for each pulse so that the pressure pulse amplitude incident on the contrast agent bubbles at their location in the absorbing tissue is practically equal for each sub range.
  • the transmitted center frequency f 0 can be less than 1MHz in the preferred embodiment .
  • a backing mount of the transducer with characteristic acoustic impedance less 30% of that of the active electro-acoustic layer may be used.
  • the improved sensitivity of Lhe receiving transducer assembly in the receive band may be facilitated by using a backing mount of the transducer with characteristic acoustic impedance greater than 150% of that of the active electro-acoustic layer .
  • the sensitivity of the receiving transducer assembly in the receive band is also facilitated by making the transducer assembly resonant in this band.
  • the ultrasound transducer array comprises an electro-acoustic active layer divided into several transducer elements with a front and a back face, a 1st thin electrode layer covering the front face, and a 22id thin electrode layer covering the back face .
  • the electrodes are electrically connected to electric terminals for coupling of energy between the electric terminals and acoustic vibrations in the transducer b elements.
  • a substrate layer is mounted on the back side of the acoustic layer with approximately the same acoustic properties as the active laye.r.
  • the back layer is mounted on an acoustically absorbing backing with 0 acoustic impedance much lower than the two layers.
  • the ultrasound transducer array further comprises at least one acoustic matching layer mounted on the front face of the active layer and acoustically in contact with the tissue.
  • the acoustic properties and 5 thicknesses of the matching layers are adjusted to facilitate improved acoustic power transfer to and from the tissue and to facilitate a wide bandwidth of the electro-acoustic transfer function to transmit a band- limited ultrasound pulse centered at f 0 into the tissue, 0 and to receive backscattered ultrasound pulses in the 3rd or 4th harmonic component, or both, of the transmit band.
  • the substrate layer also is electro-acoustically active and divided into individual transducer elements with common faces to the first transducer elements, with 5 a third, thin electrode layer on the back face of the elements, which can be combined with the 2nd or the 1st electrodes for coupling of energy between the electric terminals of the electrodes and acoustic vibrations in the transducer combined transducer elements.
  • Two of the 3 electrode layers may be connected to the transmit amplifiers to transmit the low frequency acoustic pulse, and another two of the 3 electrode layers may be coupled to the receiver amplifiers to b receive the back scattered acoustic energy from the contrast agent bubbles.
  • the method for detecting the contrast ctgent may be used for quantitating variations in tissue blood perfusion.
  • the ultrasound contrast agent in the tissue may first be destroyed uniformly with depth and direction in the tissue with a controllable degree, followed by imaging of the b backscattered signal power from contrast agent in the issue .
  • Partial destruction of the contrast agent may be performed so that the ampl tude of the backscattered signal in the 3rd or 4th harmonic component of the 0 transmit frequency band gives a regional grading of the perfusion.
  • Separate destruction pulses may be used to controllably destroy the contrast agent uniformly over the whole image field, 5
  • the contrast agent may also be fully destroyed in the tissue, and imaging may be performed at a time interval after this destruction, so that the amplitude of the back-scattered signal in the 3rd or 4th harmonic component of the .transmit frequency band gives a 0 regional grading of the refilling time of blood into the tissue, and hence the blood perfusion through the tissue.
  • the timing of the contrast agent destruction may be derived from the electrocardiogram (ECG) , and 5 imaging may be performed at a selected period in the cardiac cycle derived from the ECG.
  • ECG electrocardiogram
  • Fig. la is a graph showing the transmitted and the propagation distorted pressure pulse and Fig. lb shows the Fourier amplitude spectrum of a typically distorted pulse in tissue ,-
  • Fig. 2a is a graph showing a typical, distorted back-scattered pulse from a contrast agent bubble and Fig. 2b shows the Fourier amplitude spectrum of the pulse;
  • Fig, 3 is a block diagram of an instrument for real time imaging of contrast agent using higher harmonic components in the scattered signal for the detecti.on and imaging of the contrast agent
  • Fig. 4a is a cross section of a layered ultrasound transducer design with low damping due to low backing impedance, that can transmit an ultrasound pressure pulse with limited frequency band around f 0 , and receive back scattered signal in a frequency band around the 3f 0 and/or 4f 0
  • Fig. 4b shows a typical lateral division of the active layer into array elements in a two-dimensional matrix (the transducer is also useful for l aL and 2 nd harmonic imaging) ;
  • Fig. 5a shows the frequency transfer function from electric drive voltage to front face vibration velocity for one selection of electrodes of the transducer in Figure 4 and Fig. ⁇ b shows the same transfer function for another selection of electrodes;
  • Fig. 6a is a cross section of another layered transducer design with low damping due to high backing impedance, being able to transmit an ultrasound pressure pulse with limited frequency band around f lake, and receive back scattered signal in a frequency band around 3f 0 and/or 4f n>
  • Fig. 6b shows the frequency transfer function from electric drive voltage to front face vibration velocity for one selection of electrodes of the transducer (the transducer is also useful for l' il " and 2 ⁇ 1 5 harmonic imaging)
  • Fig. 6 ⁇ shows the frequency transfer function from electric drive voltage to the front face vibration velocity for another selection of electrodes in the transducer;
  • Fig. 7 is a block diagram of a signal 10 generator and power amplifier together with an optional filter, for electric driving of the transducer element with a band-limited voltage oscillation;
  • Figs. 8a and 8b are schematic depictions showing adequate selection of transmit amplitudes, foci, 1.5 and apertures of transmit pulses for several sub ranges can be used to obtain practically constant incident pulse amplitude and width for a large depth range;
  • Fig. 9 is a schematic depiction of how the inflow from one artery branches into capillaries that 20 gives the local blood perfusion through the tissue, and finally converges into a vein;
  • Fig. 10 shows how the filling curves of contrast agent concentration into a tissue region depend on the blood perfusion through the region.
  • the stiffness of a soft tissue increases when the tissue is compressed by the positive pressure in a transmitted ultrasound pulse.
  • the Figure displays a first harmonic band 104 with bandwidth B and centered around the center frequency f 0 of the transmitted pulse.
  • a 2" d harmonic band 105 centered around 2f 0 has lower amplitude than the 1 3C harmonic band, and a 3 r ⁇ harmonic band 105 centered around 3f n is barely visible above the noise level.
  • the 4 ch harmonic band disappears below the noise.
  • the distortion also produces sub harmonic frequency components from the time derivative of the pulse envelope.
  • this sub harmonic frequency band is centered at approximately f ⁇ /2.
  • the 2 nd harmonic component of the transmitted frequency band in the back scattered signal from soft tissue ste s from the non-linear, forward propagation distortion of the incident pulse, and linear back scattering of the distorted pulse.
  • the amount of 2 nd harmonic component in the back-scattered signal from the tissue inhomogene.ities therefore, depends on the forward propagation distortion of the incident pulse.
  • the situation is different, in that considerable non-linear distortion of the incident pulse occurs in the scattering process itself.
  • the bubble is much smaller than the acoustic wavelength ( bubble diameter ⁇ 5 ⁇ m, acoustic wavelength ⁇ 50 ⁇ m)
  • the bubble volume can expand with mainly shear deformation of the surrounding tissue (i.e. limited volume change) .
  • This phenomenon produces the following effects on volume compression/expansion of the bubble and the nearest surrounding tissue: i)
  • the volume compressibility of the bubble with nearest surrounding tissue is determined mainly by the compressibility of the bubble, where a co- oscillating volume of tissue approximately 3 times the bubble volume moves in the compression, with mainly shear deformation.
  • the bubble has a highly non-linear elasticity given by the adiabatic gas relation and the non-linear elasticity of the bubble shell. This means that a negative acoustic pressure will give a relatively large increase in the bubble diameter, as the gas pressure does not become negative.
  • the pulse pressure starts to swing positive, the mass of the surrounding tissue gets ⁇ high inward momentum, which interacts with the bubble elasticity to generate a high, non-linear increase in the pressure.
  • the non-linearity of point ii) is the basic mechanism for harmonic generation in the scattering of ultrasound from contrast agent bubbles.
  • the ga ⁇ and bubble diameter By proper selection of the ga ⁇ and bubble diameter, one can use the. resonance effect in point i) to enhance the scattering of a selected harmonic component of the incident frequency band.
  • the bubble resonance frequency is slightly larger than the incident ultrasound frequency, large non-linear scattering of the incident ultrasound pulse occurs, with a typical back- scattered pulse illustrated in Figure 2a as 201.
  • the amplitudes of the 2" d harmonic component from the tissue and from the contrast agent with venous injection from contrast agent are comparable.
  • the tissue signal appears as a background noise in the 2 u ⁇ harmonic band.
  • FIG. 3 A block diagram of an instrument for real time implementation of the contrast agent detection method, is illustrated in Figure 3.
  • the transmit beam former generates appropriate electric drive signals for the transducer array elements to steer with selectable transmit aperture both the direction, the focus, and the amplitude of a transmitted ultrasound pulse into soft tissue 302 that contains ultrasound contrast agent bubbles.
  • the transmit beam former and transmit/receive switch is steered by the control unit 305, which after the pulse transmission switches 303 to connect the transducer array to a receiver beam former 306.
  • the receiver beam former produces with range dynamic focussing the received ultrasound radio frequency (RF) signal 307 for a spatially directive receiver beam, or parts thereof, according to well known principles.
  • the received RF-signal is then fed to a filter unit 308, which in its output 309 selects frequency bands of the received signal around harmonics of the transmitted center frequency.
  • the amplitude of 309 is fed to a scan converter and display unit 311. This unit displays according to well known principles, the amplitude of 309 as an M-mode, a two-dimensional or a three dimensional image depending on the ultrasound beam scan modes of the data collection.
  • the second, third or fourth harmonic component of the transmit frequency band can be used tor the detection and imaging of the contrast agent .
  • the filtered RF- slgnal 309 can be further processed in an advanced processing unit 310 to determine for example the temporal variation of the signal intensity to monitor refilling times of blood into a region.
  • the filtered RF- taignal 309 can undergo Doppler processing to determine the velocity of the contrast agent.
  • Doppler processing can for example be used to suppress the contrast agent signal from the blood pools in the heart cavities in the periods of the heart cycle where the blood in the heart cavities moves faster than the myocardium. This can be used to enhance the display of the scattered contrast agent signal in the myocardium above the signal from the heart cavities, as described below.
  • the Doppler processing can also be used to determine rnyocardial strain from gradients in the rnyocardial velocity or the phase of the RF-signal between consecutive pulses.
  • the transducer array assembly 301 in Figure 3 must be able to transmit ultrasound frequencies in a limited frequency band B centered around f 0 , and receive the back-scattered signal with frequency components in a frequency band around 3f district and/or 4f 0 with adequate sensitivity. Damping of the transducer in the transmit band around f 0 is tolerable within limits, as adequate amplitude of the transmitted pressure pulse can be achieved by increasing the electric drive voltage of the transducer. However, damping of the transducer in the receive band is highly disadvantageous, as it irreversibly reduces the detectability of contrast agent bubbles in the tissue.
  • the source strength of the scattering bubbles is limited.
  • there is a maximal amplitude of the received contrast agent signal in the 3 rd and 4 l harmonic band and it is therefore especially important that the internal damping (power absorption) of the transducer in the receive band is low.
  • Such absorption directly drains power from the received signal, reducing the receive sensitivity and limiting the detectability of contrast agent bubbles in the tissue.
  • the transducer is a layered structure, composed of an electro-acoustic active layer 401 with a front electrode, 402, and a back electrode, 403.
  • the layer is laterally divided into a number of transducer elements, 410, where Figure 4b illustrates a possible lateral division for a two-dimensional array .
  • the elements have individual electrodes on at least one of the faces, so that the voltage across each element can be controlled individually. It is then convenient that the front ? Q electrode 402 is a ground layer, common for all elements, while the back electrode 403 is divided into individual electrodes for each transducer element. With this Selection, the front ground electrode will provide shielding of the active elements against external electromagnetic interference.
  • the active layer 401 is mounted on an acoustic substrate layer 404, with similar characteristic acoustic impedance as the active layer.
  • This substrate layer will influence transducer resonances both in the transmit and the receive bands, and is mounted on a backing material 405 with as low characteristic acoustic impedance as possible, while still giving mechanical support for the transducer. Reducing the characteristic impedance of the backing, reduces the acoustic power that is transferred into the backing and hence the transducer power losses, improving the transducer sensitivity.
  • the added acoustic layer 404 is mounted in front of the active layer 401, still having a similar characteristic impedance as the active layer, which together forms a composite layer with high characteristic impedance that participates together to form resonances in the structure.
  • the substrate layer is made of the same electro-acoustic active material as the active layer 401, and the back face of each element is coated with an electrode 409, the use of which is discussed below.
  • a set of acoustic impedance matching layers which couples acoustic energy from the active layer 401 to the tissue 406.
  • a typical selection of parameters for the transducer is to use a ceramic-epoxy composite for the acousto-active layer, which is made according to woll known principles.
  • One can then typically obtain a b characteristic impedance of the active layer of Z ⁇ 10 MRayl.
  • the thickness of the active layer is chosen so that it has a 0 short circuit 0.19 ⁇ at 3MHz .
  • the substrate layer has a characteristic impedance of 10 MRayl, and is 0,75 ⁇ thick at 3MHz.
  • the backing layer is assumed acoustically infinite, i.e. reflected waves can be neglected due to internal absorption, with a characteristic impedance of 5 O.B Rayl.
  • the matching layer 407 has a characteristic impedance of 4.3 MRayl and a thickness of 0.37 ⁇ at 3 MH,z, while the acoustic matching layer 408 has a characteristic impedance of 2.33MRayl and a thickness of 0.55 ⁇ at 3MHz.
  • the transfer function has a dip around 1.6MHz, which attenuates both transmission and reception of the 2 nd harmonic band around 2£ 0 . 0
  • a modification of this transducer design that has improved transmit efficiency, is to use the active electro-acoustic material also for the substrate layer with the added electrode 409, individually divided for each transducer element. Applying the transmit voltage 5 between electrodes 409 and 402, produces a frequency transfer function from electric voltage to front face vibration velocity shown as 502 in Figure 5b . This coupling improves the transmit efficiency at 1MHz , and reduces the transmit efficiency for the 3" 1 harmonic component of the transmit voltage at 3MHz .
  • Electrodes can also b used to transmit ultrasound pulses with frequencies in the range 2 3.4MHz and receive frequencies in the same range for I'" harmonic imaging.
  • the structure can also be used to transmit pulses with center frequency around 1.35MHz and receive 2 ad harmonic frequencies around 2.7MHz.
  • the layer parameters can be varied for further improvement of the transducer performance.
  • the inventive aspect of the design is that adequate transfer functions of the assembly is obtained with minimal damping of the transducer. This is achieved with minimizing the characteristic Impedance of the backing, and utilizing a substrate layer, to provide transducer vibration resonances both in the transmit and the receive bands.
  • the transfer bands of Figures 5a and 5b can be moved proportionally up or down in frequency.
  • a low damped transducer can also be obtained by using a backing material with characteristic impedance much higher then the impedance of the electro- acoustic active layer.
  • a design using this principle is shown in Figure 6a, where 601- shows the active electro- acoustic layer with a front electrode 602 and a back electrode 603.
  • the active layer is mounted on a substrate layer 604 with close to the same characteristic impedance as the active layer, and the substrate layer is mounted on the high impedance backing material 605, sufficiently absorbing that reflections from the far end of the backing can be neglected.
  • the b Figure shows the use of a single matching layer 606 for coupling of the acoustic power to the tissue 607.
  • the "substrate" layer 604 could al ⁇ o with the high impedance backing be mounted in front of the active layer, to form a composite layer with high acoustic impedance that gives a composite contribution to the 5 resonances of the transducer .
  • FIG. 6b shows the frequency transfer function 609 of the voltage to acoustic vibration velocity on the front face, coupling the electric voltage between electrodes 602 and 603.
  • Characteristic b impedance of the backing material 20MRayl
  • Characteristic impedance and thickness of the substrate layer lOMRayl and 0,15 ⁇ at 3MHz
  • Thickness of the active layer 0,27 ⁇ at 3MHz
  • Characteristic impedance and thickness of the matching layer 5MRayl and 0,375 ⁇ at 0 3MHz.
  • the design in Figure 6a has more internal power losses (damping) than the designs in Figures 4a and 5a.
  • Increasing the backing impedance above 20MRayl reduces the absorption, and presents challenging requirements for acoustic material development .
  • the non-linear scattering from contrast agent is more sensitive to the amplitude of the incident pulse than the linear tissue scattering.
  • incident amplitude of the transmit beam which according to the invention is obtained by matching the transmit aperture and the transmit focus to the actual absorption in the tissue, so that the amplitude of the incident pressure pulse is practically constant at all ranges.
  • One method, accordiny to the invention, to obtain such a matching is to start with preset transmit beam aperture and focussing according to a preset and selectable absorption per unit depth and frequency. The. depth variable receiver gain is then adjusted so that the 1 at harmonic scattered signal from the tissue and contrast agent is constant with depth.
  • the absorption is approximately proportional to the frequency, which gives a first harmonic gain level g x (z)
  • a(z) is the absorptioii per unit frequency and depth
  • G x is a gain level
  • the gain variation g n (z) is ideally set to
  • Figure 8a shows the axial amplitude of the three transmi pulses, while Figure 0b shows approximate width of the transmit beams in the 3 zones.
  • a circular transmit aperture is used.
  • the first zone 801 is imaged with a transmit aperture diameter of 17mm with the transmit focus at 185mm, giving the amplitude variation 804 with depth.
  • the second zone 802 is imaged with a transmit aperture diameter of 26mm with the transmit focus at 202mm, giving the amplitude variation 805 with depth.
  • the third zone 803 is imaged with a transmit aperture diameter of 38mm with the transmit focus at 265mm, giving the amplitude variation 806 with depth. Approximate width of the beams in the zones are indicated with the lines 807 for the l ',c zone, 808 for the 2 M zone, and 809 for the 3" 1 zone.
  • the beam for the 3 rrt zone has close to constant amplitude for the whole image range, but it gets a large width at low ranges. This can produce problems with interference from the contrast agent signal from the blood pools in the heart cavities, in the imaging of contrast agent in the myocardium. It is in such a situation an advantage to use the narrower transmit beams for imaging in the other zones.
  • the nea -range width of the transmit beam is not a problem, one can use a single transmit beam for imaging the whole range.
  • FIG. 9 The perfusion of blood through a tissue is illustrated in Figure 9, where 901 shows a larger vessel that feeds the region of the tissue with blood.
  • the inflow vessel branches into a capillary system 902 through which the tissue is fed with oxygen and nutrition, and metabolic byproducts are removed.
  • the capillary system converges into the venous system 903 that carries the blood away from the tissue.
  • the arterial inflow of blood is a source, and the venous outflow is a sink of contrast agent into the tissue.
  • the arterial and the venous blood flow is normally the same, and with no other sources or sinks, the concentration of contrast agent .in the tissue is in the stationary situation is a product of the volume concentration of blood in the tissue, and the concentration of contrast agent bubbles in the inflowing blood.
  • the perfusion through the tissue describes the volume of blood that flows through a unit volume of tissue per unit time, and is usually measured in s "1 .
  • the concentration of contrast agent in the tissue is in the stationary situation practically independent of the perfusion through the tissue, until close to complete blockage in the inflow vessel occurs.
  • perfusion is limited by a stenoses in a coronary artery, the blood volume in the tissue will even increase due to increased diameter of the resistance vessels.
  • the power of the back- scattered signal from a sample volume will be proportional to the number of contrast agent bubbles in the volume.
  • the signal power is therefore proportional to the product of the concentration of contrast agent bubbles in the volume and the size of the sample volume, which is determined by the length of the received pulse from a bubble, and the width of the combined transmit/receive beam.
  • the signal power will be little influenced by 20 the perfusion through the tissue until close to complete blockage of the perfusion occurs.
  • the filling time will increase with low perfusion, and the signal amplitude will reflect the perfusion rate, as shown below.
  • Similar effects can also be obtained by partial or complete destruction of the contrast agent with incident ultrasound pulses, where the degree of destruction depends on the amplitude, frequency, pulse length, and other parameters of the pulse.
  • Such pulse destruction introduces an extra Bink of contrast agent in the tissue, where the contrast agent concentration in the tissue, C (#bubbles/ml) , will depend on the perfxision rate p ( s ⁇ l ) , the contrast agent concentration in the inflowing blood, C i n (#bubbles/ml) , the blood concentration in the tissue, C b (ml/ml) and the destruction rate q (s '1 ) .
  • the relationship can be expressed by the following differential equation
  • the inflow time constant is 1/p, and for high perfusion rate we get a refilling curve of the contrast b agent concentxdtion C(t) as 1001 in Figure 10, compared to the refilling curve 1002 for low perfusion rate.
  • ECG ECG
  • Partial bubble destruction by the transmitted pulses can also be used to further enhance the contras 0 agent signal above the tissue signal, by transmitting several pulses in the same beam direction.
  • the bubble destruction will then introduce a reduced correlation time of the contrast agent signal compared to the tissue signal .
  • the tissue signal can then be relatively b attenuated by high-pass filtering of the back-scattered signal along the pulse number coordinate.
  • the relative blood volume in the myocardium is the relative blood volume in the myocardium.
  • the signal power from the myocardium is hence ⁇ - 0 12dB of the signal power from the blood pools in the heart cavities.
  • Sidelobes of the image beam together with limited range resolution in the ultrasound image hence, will give a corroboration of the signal from the heart cavities into the signal from the myocardium.
  • This b gives a lower threshold for the rnyocardial contrast agent that can be discriminated in the rnyocardial regions that are corroborated by the signal from the blood .
  • the blood in the heart cavities at times during the cardiac cycle is moving much faster than the blood in the myocardium.
  • One example of such processing is to average the received radio frequency signal at each range and beam direction for N pulses, which will attenuate signal components from scatterers with high velocity, compared to those from scatterers with lower velocity.
  • This processing can be done in the Advanced processing unit 3.10 of Figure 3.
  • the backscattered pulse from the contrast agent bubble is as short as possible. This is achieved through two mechanisms: 1) The transmitted pulse centered around f detox must be as short as possible, while limiting the bandwidth of significant frequency components in the pulse. 2) For a short incident pulse, the length of the scattered pulse will depend on the polarity of the incident pressure pulse. The reason for this is that the bubble produces a particularly strong, non-linear and resonant oscillation in the scattered pulse at the turn of the radius oscillation at its minimum. This occurs shortly after the incident pressure pulse swings from negative to positive pressure. This non-linear oscillation is tho main source of non-linear distortion in the scattering from contrast agent bubbles. For short incident pulses, the number of significant negative to positive pressure swings in the incident pulse depends on the polarity of the pulse, and the length of the 2nd and higher harmonic components in the scattered pulse hence depends on the polarity of the incident pulse.

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Abstract

La présente invention concerne une méthode de détection d'un agent de contraste ultrasonore dans un tissu mou et de quantification de la perfusion sanguine à travers des régions tissulaires par le biais de la détection de l'agent de contraste dans le tissu.
PCT/NO2001/000022 2001-01-19 2001-01-22 Procede de detection d'un agent contraste ultrasonore dans un tissu mou, et de quantification de la perfusion sanguine a travers des regions tissulaires Ceased WO2002056666A2 (fr)

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WO2004110279A1 (fr) * 2003-06-12 2004-12-23 Bracco Research Sa Estimation de debit sanguin par ajustement de courbe de regeneration dans l'imagerie de contraste par ultrasons
EP1674038A1 (fr) * 2004-12-23 2006-06-28 Bracco Research S.A. Système pour extraire des informations morphologiques grâce à un procédé d'évaluation de perfusion
US8021303B2 (en) 2003-06-12 2011-09-20 Bracco Research Sa System for extracting morphological information through a perfusion assessment process
US8036437B2 (en) * 2005-04-14 2011-10-11 Bracco Research Sa Perfusion assessment method and system based on animated perfusion imaging
US8496591B2 (en) 2004-12-23 2013-07-30 Bracco Suisse S.A. Perfusion assessment method and system based on bolus administration
EP3829421A4 (fr) * 2018-07-30 2022-04-20 The Regents of The University of California Systèmes et procédés de quantification de flux sur la base d'images intra-intervention

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JP3015481B2 (ja) * 1990-03-28 2000-03-06 株式会社東芝 超音波プローブ・システム
US5410205A (en) * 1993-02-11 1995-04-25 Hewlett-Packard Company Ultrasonic transducer having two or more resonance frequencies
US5724976A (en) * 1994-12-28 1998-03-10 Kabushiki Kaisha Toshiba Ultrasound imaging preferable to ultrasound contrast echography
EP1374777A1 (fr) * 1995-10-10 2004-01-02 Advanced Technology Laboratories, Inc. Imagerie diagnostique à ultrasons utilisant des agents de contraste
GB9708246D0 (en) * 1997-04-24 1997-06-18 Nycomed Imaging As Improvements in or relating to ultrasound imaging
US5980459A (en) * 1998-03-31 1999-11-09 General Electric Company Ultrasound imaging using coded excitation on transmit and selective filtering of fundamental and (sub)harmonic signals on receive
US5957852A (en) * 1998-06-02 1999-09-28 Acuson Corporation Ultrasonic harmonic imaging system and method

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WO2004110279A1 (fr) * 2003-06-12 2004-12-23 Bracco Research Sa Estimation de debit sanguin par ajustement de courbe de regeneration dans l'imagerie de contraste par ultrasons
JP2006527041A (ja) * 2003-06-12 2006-11-30 ブラッコ・リサーチ・ソシエテ・アノニム 超音波造影画像において補充曲線フィッティングを用いる血流評価法
KR101025490B1 (ko) * 2003-06-12 2011-04-04 브라코 인터내셔날 비.브이. 초음파 콘트라스트 조영에서 보충 커브 피팅을 통한 혈류 개산
US8021303B2 (en) 2003-06-12 2011-09-20 Bracco Research Sa System for extracting morphological information through a perfusion assessment process
US8491482B2 (en) 2003-06-12 2013-07-23 Bracco Suisse S.A. Blood flow estimates through replenishment curve fitting in ultrasound contrast imaging
NO337721B1 (no) * 2003-06-12 2016-06-13 Bracco Suisse Sa Blodstrømningsestimater ved etterfyllings-kurvetilpasning i ultralydkontrastavbildning
EP1674038A1 (fr) * 2004-12-23 2006-06-28 Bracco Research S.A. Système pour extraire des informations morphologiques grâce à un procédé d'évaluation de perfusion
WO2006067203A1 (fr) * 2004-12-23 2006-06-29 Bracco Research Sa Systeme d'extraction d'une information morphologique par un procede d'evaluation de perfusion
US8496591B2 (en) 2004-12-23 2013-07-30 Bracco Suisse S.A. Perfusion assessment method and system based on bolus administration
US8036437B2 (en) * 2005-04-14 2011-10-11 Bracco Research Sa Perfusion assessment method and system based on animated perfusion imaging
EP3829421A4 (fr) * 2018-07-30 2022-04-20 The Regents of The University of California Systèmes et procédés de quantification de flux sur la base d'images intra-intervention
US11769252B2 (en) 2018-07-30 2023-09-26 The Regents Of The University Of California Systems and methods for intra-procedure image-based flow quantification

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