RECURSIVE ULTRASOUND IMAGING

Abstract: A method and an apparatus for recursive ultrasound imaging is presented. The method uses a Synthetic Transmit Aperture, but unlike previous approaches a new frame is created at every pulse emission. In receive, parallel beam forming is implemented. The beam formed RF data is added to the previously created RF lines. To keep the level of the signal, the RF data obtained previously, when emitting with the same element is subtracted from the RF lines. Up to 5000 frames/sec can be achieved for a tissue depth of 15 cm with a speed of sound of c = 1540 m/s. The high frame rate makes continuous imaging data possible, which can significantly enhance flow imaging. A point spread function 2 DEG wide at -6 dB and grating lobes of </= -50 dB is obtained with a 64 elements phased array with a central frequency ƒ0 = 3 MHz using a sparse transmit aperture using only 10 elements (Nxmt = 10) during pulse emission. The corresponding images have the quality of a dynamically focused image in transmit and receive. The dynamic focusing gives a small sampling volume and the capability to view small blood vessels and obtain the velocity profiles within the vessels with lower variance than with normal imaging.

Description: Recursive Ultrasound Imaging Field of the invention The invention relates to an apparatus for fast imaging of eg the human body using ultrasound The image is made using a multi-element probe in which all or selected ones of the elements are used as transmitters The reflected signal from the object is then measured by all elements and ultrasound beams are simultaneously focused throughout the imaging region The image is updated with the new information every time a new element or a group of elements are used as a transmitting aperture Hereby the image is continuously updated and can be used for probing moving structures and enhance the imaging of blood velocities Background of the invention The technological development has constantly led to improvements in ultrasound beam formers as described in [1] The introduction of digital beam formers made the dynamic receive focusing possible Unfortunately it is not possible to dynamically control the transmitted acoustic energy, and only a fixed focus is used in transmit The basic concept of focusing is to constructively add energy in the combined transmit-receive signal Ideally, by means of the receive processing it is possible to compensate for the phase changes in the transmit This corresponds to a dynamically focused transmit and receive imaging system A composite image (obtained by multiple emissions, focused at different depths in transmit) is the ideal reference image One way to obtain a dynamic transmit focus is to use synthetic array imaging There are three variations of the synthetic aperture imaging 1 using a synthetic transmit aperture, 2 using a synthetic receive aperture, and 3 using a combination of synthetic transmit-receive aperture All of these combinations have been studied, and have shown different advantages and drawbacks The time necessary to acquire a single image T αcquιre mαge is proportional to the number of emissions N em , ss ,or, , the time necessary to record the reflected ultrasound wave from a single emission T rec0 r d emission , the number of scan lines N , and it is inversely proportional to the number of the parallel receive beam formers N para ιiei T acquireimage = T acquirescanhne ■ N emission - N 1 IN parallel ( VΩ The acquisition time for an image, with a typical depth of 15 cm assuming that the speed of sound is 1540 m/s, is 200 μs. IfN emss , 0 „ - 64 and N = N para ιi e i then acquire i ge = 12.8 ins, which results in 78 frames/sec. For every new frame in the imaging process, the data acquired for the previous frame is discarded. The recursive imaging method and apparatus according to the invention uses the beam formed lines from the previous frame to create a new frame after every emission. This results in N emrø , 0 „ = 1, T acquir image = 200 μs, and a frame rate of 5000 frames/sec. The invention uses a synthetic transmit aperture and preferably receives with the full aperture. The prior art synthetic transmit aperture focusing [2], [3] is presented below. Based on it, the new recursive ultrasound imaging technique according to the invention is mathematically derived and discussed below. One of the problems in emitting with only one element is the signal-to-noise ratio, since there are physical limitations to what power can be sent with a single transducer array element. The problem was previously studied [2], [4], and a solution using multiple elements in transmit was suggested in [2]. This problem is discussed below. The other problem is the presence of motion artifacts due to the time of acquisition for a single image. This is avoided with the invention by decreasing the number of emissions necessary to create one frame. This can be done with a sparse synthetic aperture, and the results of the decreased number of emissions is presented. Prior art synthetic array imaging Phased linear arrays are used for creating sector B-mode images as shown on Fig. 2. The image consists of a number of scan lines N. The scan lines / = (1..N) have a common origin. Each of them has a different angle θj with the normal vector to the transducer's surface For simplicity all considerations are made only in the z-x plane, and it is assumed that a single transducer element emits a cylindrical wave propagating at a constant speed c in a linear medium, as shown in Fig 3 The measurement situation is shown in Fig 4 Element i with center coordinates (x„z,) emits a spherical wave The wave front reaches point P(x P ,z P ) after time where c is the speed of sound If all the elements are excited, they will form a pressure field in front of the transducer, which is a sum of the fields created by each of the elements The emitted waves reach point P at different time instances depending on the element positions In order to align the wave fronts and thereby focus the acoustic energy in point R, the signals from the individual array elements must be appropriately delayed In the calculation of the delays, one point from the transducer is selected as a reference point All the delays are calculated relative to the time necessary for a sound wave to travel the distance between the reference point and the focal point For a sector image, the center of the transducer array is the usual choice The reference point is the center of element C(x c ,z c ) for the case depicted on Fig 4 The delay for element i for focusing the energy at point P(x p ,y p ) is calculated by " i = *CP ~ *iP d. = ( c - χ pΫ + ( ' c - Z p Ϋ - ΛI X , - X p) 2 + (z, - -pΫ (3) Fig 5 shows the geometry used for the simulations The origin of the coordinate system O(0,0) lies in the middle of the physical linear array transducer All scan-lines (/ = 1 Ni ) start from the origin O and have angle θj with the z-axis The focal points lie on these scan lines and have coordinates x fl = r f ■ sin Θ, (4) z β - r f cosΘ, where ; is the axial distance to the focal point. Dynamic focusing is obtained if r/ changes in time as r f = ck/f s (5) where k is the sample number and f s is the sampling frequency. In conventional imaging systems the transmission uses a fixed r/ , whereas the receive can be dynamically focused along the current scan line One way to create a synthetic aperture image is to emit with all the elements from the transducer array one at a time (see Fig. 6). During reception data is recorded and beam formed. After a number of emissions all the beam formed RF lines from the separate emissions are summed to create the scan-lines that are envelope-detected and displayed. The beam formed signal from emission n with element / for line / is s^ { t . The number for the emission n (0 < n < ∞) is relative to the beginning of the continuous imaging process. This number is relevant only for the recursive imaging and in this section a simplified notation is used: s h (t) - . The time t is relative time from the emission of the pulse. The RF signal that is envelope-detected and displayed is: ,(>)= ∑A( (6) s h (t) = ∑a h] - s lJ (t - d hj ) (1) where N^, is the number of transmit elements, N rcv is the number of receive elements, s, t) is the recorded data, ; v is a weighting coefficient (apodization), and d h} is the delay for image line / , when emitting with element i and receiving with element j. The delays dj y is the sum of the delays for elements / ' andj calculated with formula (3) for the points in line / For the case in Fig. 5, the delays are calculated by: ^( χ , - χ fl ) 2 + - fl 2 + z. l( x j - χ β ) ~ + z, , = (8) This is the basic formula for a geometrical quadrature focusing. In Eq. (8) xp , Zβ) are the coordinates of the focal point, x, is the coordinate of the center of the th element and x, is the coordinate of the center of the t'th element. In transmit, the delay for element / is equal to zero, and in receive the delay for element / is equal to dι v . The synthetic array imaging assumes that the tissue under investigation is stationary, which is often not the case in medical ultrasound. The images will, thus, be blurred and the synthetic array imaging cannot be used for velocity estimation. Summary of the invention It is the object of the invention to overcome this deficiency and disadvantage of the known kinds of apparatus With the invention this object is achieved by an apparatus that continually updates the image using the current pulse emission thereby obtaining a continuous image for tracking moving structures. The concept of the invention is as follows: A new frame n is created, based on the previous frame n-1 and the data acquired at the last emission. The calculation procedure is derived from a synthetic array aperture imaging technique as presented and further developed. Recursive imaging with the invention A simple example for synthetic imaging is shown in Fig. 6. Here n is the number for the current emission and Nχm t is the number of transmissions between two emissions with the same element. is the beam formed signal at emission n with element k for line / . The relation between n and k is given by: £(«) = (« mod N™,) + 1 (9) The first image can be obtained after N^, emissions. Equation (6) can be rewritten as: Equation ( 10) shows that a new frame can be formed after any emission n ≥ Nχ mt by summing the beam formed lines s <m) h(m) , n -Nxmt < m ≤ n . For two consecutive emissions the expression is: m-(n §-Nx«m«t).> ■ o en) ύ j - d, = a k{n) ■ s Ik(n) (t) - a k(n) ■ s lk{n _ Nxmt) (t) (12) θ(«) _ (n-l) , „ . „(«) (f _ n (n-Nxmt) (f j — , ^ u k( ) ύ lk(n)\' I U k(n-Nxmt) ύ lk(n-Nxmt) \ l ) From equation (9) it can be seen that k(n) = k(n -N xm t). In recursive imaging a new frame is created at emission n by adding the new information Info (n) to the image Image ( " ~!) and subtracting the information obtained at emission n Info <n'Nxm,) . The number of summations per sample is decreased rrom , in Eq. (6) to only two. Since a new frame consisting of a number of simultaneously beam formed lines is created at every pulse emission, the image is updated after every 200 μs when c = 1540 m/s, and the data is acquired to a depth of 15 cm. Add-only recursive imaging To decrease the amount of the necessary storage memory the calculation procedure can be modified to the add-only recursive imaging Consider the following equations R = ∑4 (13) 1=1 -4, = A - const => B = N - A = N - A + A - A = ^ - N - A + A = x^- - B + A (14) N N The above equations can be used to derive another formula for the recursive imaging -S / ( " ( = c, - -S / ( ^ ) + c 2 . ">( (15) where Cj and c 2 are weighting coefficients The constants must be chosen to ensure a constant signal level and the following condition must be fulfilled S ; (n) (t) = S, ( " _1) (t) The difference between equations (12) and (15), is that instead of being subtracted the information obtained by emitting with element k decays exponentially with time In this way the information from the past is less prone to introduce motion artifacts in the image The other benefit is that less memory is needed, since only two frames are stored Equation (15) can be rewritten as a sum of the signals from two consecutive emissions S, ( ">(t) = ∑b"- m - s m) t) + sil ) (t) (16) assuming that c 2 = 1 and Cι = b < 1 Nχm t < n < ∞ is the number of the frame, which is equal to the number of the current emission The same element is used after every Nχm t emissions Equation (16) can be expressed as a sum of the contributions from the emissions with the different elements Here contribution of one element is the amount of information C ° , added to the signal S ; (π) by the emissions with element k up to the moment of emission n . The contribution of the current emitting element (k(ri) = i) to S n is: C n) (t) = + b Nxn " ■ S { - Nxm,) (t) + b Wxmt ■ s "- 2Nxm,) (t) + ... (18) This is a geometrical progression. If the tissue is motionless then: If b = 0.8 and N™, = 64 then . This means that at emission n the contribution from the emitting element is approximately equal to the newly acquired information. With the chosen values of b and N ^ , the contribution to the signal Si from the previous emission with the same element is 128 dB lower than the contribution from the current one and motion artifacts can be neglected. Improving the signal-to-noise ratio The advantage of the above presented imaging approach is that a dynamically focused image in transmit and receive is obtained. This is possible because a single small transducer element emits an almost spherical wave, which unambiguously determines the propagation time of the ultrasound energy. The drawback is that the energy, which can be sent into the body with only one array element is not enough for obtaining a high penetration depth. One way to increase the penetration depth is to emit energy with several elements as suggested in [2]. The delays of the elements are set in order to approximate the radiation pattern of a single element. Here the delays are calculated by: where a is the index of the central element, x a is the coordinate of the central element of the active aperture, x, a is the coordinate of the element whose delay in transmit d ιa is calculated, and R^,, is the radius of the spherical wave. If the number of elements in the active aperture is N actιve , then R^, r is selected to fulfill the condition: R χm , ≥ (N actn , e - )l2 - pitch (21) where pitch is the distance between the centers of two neighbor elements in the array. The elements' indices are relative to the index of the central element and are given by: ia = -(Nacve - 1 )/2 ... (N fl -.,v- - 1 )/2 (22) For convenience N act ne is usually an odd number to ensure the presence of a central element. The delays in receive are calculated by formula (8), assuming that the transmitting element has the coordinate x a . The result of using multiple elements in transmit is that the signal-to-noise-ratio is increased. Let the number of transmissions be N^ and the number of receiver elements N rcv . Then the signal-to-noise ratio is proportional to: If the number of the elements in the active sub-aperture is N actιve then the S R becomes proportional to: SNR ~ jN ctlve - N xmt - N rcv Thus, using an active aperture in transmit with N act ,ve = 11 results in an 11 dB increased in SΝR. By using this approach a continuous image consisting of a number of simultaneously beam formed lines can be made at the pulse repetition frequency of the emitted ultrasound, which makes it possible to follow tissue motion By using a array transducer a sector scan image can be made By using a matrix transducer a volumetric image can be made By using several transducer elements during the transmission an improved signal-to- noise ratio can be obtained By using recursive imaging velocity images can be improved and velocity distributions can be found throughout the image Brief description of the drawings Fig 1 shows a schematic block diagram of the main components of the system Fig 2 shows a sector image created with phased linear array transducer Fig 3 shows a radiation pattern of a single transducer element Fig 4 shows the distance traveled by the ultrasound signals from each of the transducer elements to a point P Fig 5 shows a delay calculation The reference point is O(0,0) The focal point f(x ,zj) lies on the current scan line 1 The z coordinate of the transducer elements is zero, and it is not used in the notations for simplicity Fig 6 shows the creation of a synthetic transmit aperture image The black square indicates the position of the element used for emission Fig 7 shows an approximation of the radiation pattern of a single element with a sub- aperture containing 11 elements The indices of the elements are relative to the center of the sub-aperture Fig 8 shows the excitation pulse used in the simulations Fig 9 shows an illustration of the notation in the plots of the point-spread function The number of the scan line determines the angle Fig 10 shows the relation between the number of the skipped elements between two emissions and number of emissions On this Figure a linear array with 7 transducer elements is shown The number of elements that are skipped is 2, indicating 3 emissions Fig 11 shows a point-spread function of a synthetic transmit array focusing system N™, = 64, N actιve = 1 Fig 12 shows a point-spread function of a synthetic transmit array focusing system N xmt = , N a ct,ve = 1 Fig 13 shows a point-spread function of a synthetic transmit array focusing system Nxrnt = 64, N actιve = 11 Fig 14 shows a point-spread function of a synthetic transmit array focusing system Nxmt = 64, N a ct,ve = 11 Transmit apodization is applied. Description of the preferred embodiment In Fig 1 is shown an example of a preferred embodiment of an apparatus according to the invention This embodiment of the invention has its application within diagnostic medical ultrasound A typical example is visualization of the organs in the human body or the determination of blood flow In Fig 1 the specific setup of the measuring apparatus itself is indicated schematically It comprises a generator or pulser 1 , an emit beam former 2 for selecting the transmission elements, a linear array ultrasound emitting transducer 3, a linear array ultrasound receiving transducer 5, a receive beam former 6 receiving signals from the receiving transducer 5, and an update processor for recursively updating the images The pulser 1 generates a pulsed voltage signal with sinusoidal oscillations at a frequency of 3 MHz in each pulse, that is fed to the emit beam former 2 The emit beam former 2 splits up the signal from the pulser into a plurality of signals which are being fed to one or more of the respective elements of the emitting transducer array 3 The emit beam former 2 is capable of individually attenuating and delaying the signals to each or some of the elements of the transducer array 3 The ultrasound is then reflected by the object 4 and received by the elements of the transducer array 5 All of theses signals are then combined to focus all of the beams in the image in both transmit and receive in the beam processor 6 and the simultaneously focused signals are used for updating the image in the processor 7 In the preferred embodiment the same linear array transducer is used for both emitting and receiving the pulsed ultrasound field It consists of 64 elements with an element width of 0 26 mm and a spacing between neighboring elements of 0 03 mm The height of the elements is 10 mm Simulations experimentally document the functionality of the apparatus The simulation is performed using the impulse response method developed by Tupholme and by Stepanishen [5] in the implementation developed by Jensen and Svendsen [7] The high accuracy of this approach, when compared to measurements, is described in Jensen [8] The paper showed that the simulations were within 1% of the measured ultrasound fields The simulation approach is applicable for pulsed fields and is used for three-dimensional modeling the response of multiple scatterers In the preferred embodiment individual elements are pulsed consecutively and the received signals are measured on all the elements of the transducer The ultrasound beams are then focused in both transmit and receive in all directions of the image After each emission the old information from the previous emission with the element is subtracted and the new information added The image is thereby continuously updated The simulation parameters are listed in Table 1 and Table 2 A sampling frequency of f s = 105 MHz is used for the simulations in order to get high precision in the numerical calculations of the acoustical field However, when creating the results, the signal is decimated 10 times after the beam forming The excitation pulse used is shown in Fig 8 The point spread function (psf) details the quality of the ultrasound system It was obtained by placing a point scatterer in front of the transducer, and creating a B-mode image (transmit - reflect - receive) The RF lines are envelope detected by a Hilbert transformation and logarithmically compressed The displayed images are 90 ° phased array images. The scan-conversion algorithm performs a bi-linear interpolation before displaying the images The plots of the point-spread function are created by taking the maximum value of the envelope-detected signal (see Fig 9) On the plots, RANGE denotes the maximum axial distance from the transducer, START and END determine the depth range for the point-spread functions Axial distance shows the depth at which the maximum of the signal is found Another parameter given on the figures is "Number of skipped elements" It indicates how many emissions are made before the same element is used again The relation between these parameters are given in Fig 10 The elements are always used in transmit in the same order For instance in Fig 10 this means that the emission order in time is 1, 4, 7, 1, 4, 7, This is necessary when equation (12) is used for imaging because it involves subtracting the information from the previous emission However, if equation (15) is used, the elements can "interleave" and have a different emission patterns, e g , 1, 4, 7, 2, 5, 3, 6, 1, 4, 7 The received signals, are not weighted (i e there is no apodization in transmit) Fig 11 shows the psf for 64 emissions with one element at a time Reducing the number of transmitting elements to 13 gives the pfs shown in Fig 12 The data for the psf s are shown in Table 3 In order to increase the penetration depth and the signal-to-noise ratio, more than one element can be used during transmission The idea is to send a spherical wave with eleven elements A Hanning window is applied on the amplitudes of the sending elements The received signals are not weighted (no transmit apodization is applied) Fig 13 shows the psf for 64 emissions, and Fig 14 when apodization is used during the emissions The data for the psf s are shown in Table 4. Tables System parameter Notation Value Unit Speed of sound c 1540 m/s Central frequency of excitation fo 3 MHz Central wavelength of excitation λo 0 51333 mm Sampling frequency fs 105 MHz Oscillation periods N 3 Table 1 General parameters Transducer parameter Notation Value Unit Central transducer frequency / o ( "-> 3 MHz pitch pitch 0 257 mm Width of a single element w 0 227 mm Kerf Krf 0 030 mm Height of the element h 10 mm Number of receive elements N rcv 64 Relative two-sided -6 dB bandwidth B ≥ 70 % Table 2 Parameters of the simulatec phased array transducer Number of -3 dB -40 dB Position of Grating lobe emissions bandwidth bandwidth grating lobe level 64 1 01 ° 5 03 ° - - 22 1 01 ° 5 03 ° ±40 ° -58 dB 13 1 01 ° 5 03 ° ±20 ° -54 dB 10 1 01 ° 5 03 ° ±16 ° -51 dB 8 1 01 ° 5 03 ° ±13 ° -48 dB 1 01 ° 5 03 ° ±10 ° -45 dB Table 3 The point spread function of a system with N ac tιv- = 1 Number of -3 dB -40 dB Position of Grating lobe emissions bandwidth bandwidth grating lobe level 64 1 01 ° 5 03 ° - - 13 1 01 ° 5 03 ° ±20 ° -53 dB 8 1 01 ° 5 03 ° ±13 ° -48 dB Table 4 The point spread function of a system with N actl ve = 11 References [1] K E Thomenius Evolution of ultrasound beam formers In Proc. IEEE Ultrason. Symp., pages 1615-1621, 1996 [2] G R Lockwood, J R Talman, and S S Brunke Real-time 3-d ultrasound imaging using sparse synthetic aperture beam forming IEEE Trans. Ultrason., Ferroelec, Freq. Contr., 45 980-987, 1998 [3] G R Lockwood and F S Foster Design of sparse array imaging systems In Proc. IEEE Ultrason. Symp., pages 1237-1243, 1995 [4] J T Ylitalo On the signal-to-noise ratio of a synthetic aperture ultrasound imaging method Eur. J. Ultrasound, pages 277-281, 1996 [5] G E Tupholme Generation of acoustic pulses by baffled plane pistons Mathemattk t, 16.209-224, 1969 [6] P R Stepanishen Transient radiation from pistons in an infinite planar baffle J. Acoust. Soc. Am., 49 1629-1638, 1971 [7] J A Jensen and N B Svendsen Calculation of pressure fields from arbitrarily shaped, apodized, and excited ultrasound transducers IEEE Trans. Ultrason., Ferroelec, Freq. Contr., 39.262-267, 1992 [8] J A. Jensen A model for the propagation and scattering of ultrasound in tissue J. Acoust. Soc. Am., 89 182-191, 1991a

Claims: Claims 1. An apparatus for acquiring images of an object, the apparatus comprising: a generator for generating excitation signals of multiple pulses of wave energy, and an emitting transducer with a plurality of transducer elements for transforming said excitation signals into pulses of wave energy and, using, for each pulse of wave energy, selected ones of said transducer elements, for emitting said pulses of wave energy towards said object, a plurality of receiving transducer elements for receiving, from said object, signals generated by interaction between said emitted pulses of wave energy and said object, a receive beam former for delaying and processing signals from said plurality of receiving transducer elements individually relative to each other, and a processor for processing delayed signals from said plurality of receiving transducer elements to form images of said body, where a predetermined number of processed and delayed signals are used to form an image, wherein images are updated with a number less than said predetermined number of processed and delayed signals. 2. An apparatus according to claim 1 wherein images are updated after each emission of energy. 3. An apparatus according to claim 2 wherein focused simultaneous beams are updated after each emission of energy according to =( n-Nxmt) C(«) _ O.n-1) , „ . „(") (f _ α (n-Nxmt) , f °1 — °1 ~ u k(n) Λ lk(n) ' ) U k(n-Nxmt) Λ lk{n-Nxmt) \' ) where sj- is the beam formed signal at emission n with element k for image line / , a k(m) is a weighting coefficient, and N ΛTOf is the number of transmissions between two emissions with the same element 4 An apparatus according to claim 2 wherein said focused simultaneous beams are updated after each emission of energy according to S / ( " ) (t) = c 1 . S / ( "- ,) + c 2 - " ) (t) (24) where Cj and c 2 are weighting coefficients 5 An apparatus according to claim 2 wherein said wave energy is sound energy. 6 An apparatus according to claim 2 wherein said sound energy is ultrasound energy 7 An apparatus according to claim 2 wherein said wave energy is electromagnetic energy 8 An apparatus according to claim 1 wherein a number of elements are used at the same time during transmission 9 An apparatus according to claim 8 wherein said focused beams are updated after each emission of energy 10 An apparatus according to claim 9 wherein said focused beams are updated after each emission of energy according to ςs(n) _ f-tn-1 ) , Λ") ( f _ (n-Nxmt) ^ °1 — ύ l ^ u k ( ^Ik fn) ^' / U k(n-Nxmt) ύ lk (n-Nxmt) ' ) where is the beam formed signal at emission n with element k for image line / , a k(m) is a weighting coefficient, and N ^ ,- is the number of transmissions between two emissions with the same element 11 An apparatus according to claim 9 wherein said focused beams are updated after each emission of energy according to where Cj and c 2 are weighting coefficients 12 An apparatus according to claim 9 wherein said wave energy is sound energy 13 An apparatus according to claim 9 wherein said sound energy is ultrasound energy 14 An apparatus according to claim 9 wherein said wave energy is electromagnetic energy 15 A method for acquiring images of an object, the method comprising generating excitation signals of multiple pulses of wave energy, transforming, using a transducer with a plurality of transducer elements, said excitation signals into pulses of wave energy and, using, for each pulse of wave energy, selected ones of said transducer elements, for emitting said pulses of wave energy towards said object, receiving from said object, using a transducer with a plurality of transducer elements, signals generated by interaction between said emitted pulses of wave energy and said object, delaying and processing signals from said plurality of receiving transducer elements individually relative to each other, processing delayed signals from said plurality of receiving transducer elements to form images of said body, where a predetermined number of processed and delayed signals are used to form an image, wherein images are updated with a number less than said predetermined number of processed and delayed signals 16 A method according to claim 15 wherein images are updated after each emission of energy 17 A method according to claim 16 wherein focused simultaneous beams are updated after each emission of energy according to S >(t) = ∑a t(m sff m) (t) o(n) _ (n-l) , (") (f _ a (n-Nxmt) (f °1 °l ^ U k(n) ύ lk(n) V U k(n-Nxmf) ύ lk(n-Nxmt) l ) where s l ( k n) is the beam formed signal at emission n with element k for image line / , a k(m ) is a weighting coefficient, and Nχm t is the number of transmissions between two emissions with the same element 18 A method according to claim 16 wherein focused simultaneous beams are updated after each emission of energy according to S " ) (t) (24) where C \ and c are weighting coefficients 19 A method according to claim 16 wherein said wave energy is sound energy 20 A method according to claim 16 wherein said sound energy is ultrasound energy 21 A method according to claim 16 wherein said wave energy is electromagnetic energy 22 A method according to claim 15 wherein a number of elements are used at the same time during transmission 23 A method according to claim 22 wherein said focused beams are updated after each emission of energy 24 A method according to claim 23 wherein said focused beams are updated after each emission of energy according to (n) _ (n-l) , (n) ( f \ _ „ (n-Nxmt) , f s. °1 — °1 ^ U k ) Λ lk(n) ' I k(n-Nxmt) ύ lk(n-Nxmt) \ l ) where s l ( k n) is the beam formed signal at emission n with element k for image line / , a k(m) is a weighting coefficient, and Nχm t is the number of transmissions between two emissions with the same element 25 A method according to claim 23 wherein said focused beams are updated after each emission of energy according to Sj' (t) = c, - S}^ + c 2 - ^ t) (25) where ci and c are weighting coefficients 26 A method according to claim 23 wherein said wave energy is sound energy 27 A method according to claim 23 wherein said sound energy is ultrasound energy 28 A method according to claim 23 wherein said wave energy is electromagnetic energy

Inventors: JENSEN JOERGEN ARENDT, NIKOLOV SVETOSLAV, JENSEN, JOERGEN, ARENDT, NIKOLOV, SVETOSLAV
Applicants: B K MEDICAL AS, JENSEN JOERGEN ARENDT, NIKOLOV SVETOSLAV, B-K MEDICAL A/S, JENSEN, JOERGEN, ARENDT, NIKOLOV, SVETOSLAV
Application year: 2000
Main classification: G10K11
IPC codes: G01S7, G01S15