Patent Publication Number: US-2023143880-A1

Title: Three dimensional color doppler for ultrasonic volume flow measurement

Description:
This invention relates to medical diagnostic ultrasound systems and, in particular, to ultrasound systems which produce quantified measurements of the volume flow of blood through the heart or a blood vessel. 
     Ultrasound has long been used to assess various parameters of blood flow in the heart and vascular system using the Doppler principle. The basic Doppler response is flow velocity, which can further be used to determine additional characteristics of blood flow. One characteristic of interest to cardiologists is the volume flow of blood through a vessel. Early efforts to estimate volume flow consisted of multiplying a measurement of the mean velocity of blood flow by the nominal cross-sectional area of a blood vessel. However, these early efforts had shortcomings due to the need to make certain estimates. One is that the vessel lumen is circular. Another is the estimation of the mean velocity from a single Doppler measurement or from a qualitative assessment of spectral Doppler data. Velocity measurement must also be corrected for the angle between the ultrasound beam direction and the direction of flow. Yet another consideration is the laminar flow profile in the presence of stenosis. 
     A further complication arises due to the pulsatility of arterial flow. While venous flow is substantially constant, arterial flow is constantly changing over the heart cycle. Thus, the standard techniques often lack for user independency and repeatability. Some of these demands have been eased by the advent of 3D ultrasound to assess flow conditions and particularly its ability to acquire volume blood flow information. With 3D imaging, the full vessel lumen can be imaged and a sequence of 3D image data sets acquired for later replay and diagnosis. When data of the full volumetric flow in the vessel is acquired in the data sets, the image data can be examined during post-acquisition diagnosis to assess the flow profile. Different 2D image planes can be extracted from the 3D data in multi-planar reconstruction (MPR), so that an image plane of a desired orientation through a vessel can be examined. Three dimensional imaging thus addresses many of the static imaging challenges which are problematic with 2D flow estimation. However, a substantial amount of time can be required to acquire Doppler data in three dimensions, reducing the temporal accuracy of analyzing volume flow. 
     Accordingly, it is desirable to develop more robust techniques for accurately assessing volume flow in the presence of flow pulsatility and erratic heartbeats. It is further desirable to improve the accuracy and reliability of Doppler angle measurements needed to refine the accuracy of Doppler flow velocity values. 
     In accordance with the principles of the present invention, a diagnostic ultrasound system is described which uses a 3D imaging probe to make volume flow measurements. The probe is preferably a two dimensional matrix array probe which is operated in the biplane mode. One of the imaging planes is manipulated so as to acquire a long axis view of a vessel where volume flow is to be measured. The plane of the other biplane image is aligned with the beam direction of the first image, and images the target vessel obliquely in a transverse view. The beam direction of the longitudinal view image and the flow direction seen in the long axis view thus determine the Doppler angle for Doppler angle correction of velocity values obtained from the transverse image. The Doppler data acquisition rate is relatively high, since only two planar images need to be acquired instead of a full 3D volume acquisition. And Doppler angle correction proceeds directly from the Doppler angle found for the longitudinal image. Known methods of volume flow measurement such as the Gaussian surface integral method can thus be used with high accuracy and repeatability. 
     In a method of the present invention an ultrasonic diagnostic imaging system is used to conduct an ultrasound exam to measure volume flow. Scanning is performed with an ultrasound probe adapted to operate in a biplane mode to acquire a first Doppler image of a target vessel in a long axis view. Scanning is performed with the ultrasound probe in the biplane mode to simultaneously acquire a second Doppler image in a transverse view of the target vessel in an image plane aligned with a Doppler angle of the first image. The two images are displayed simultaneously. Angle correction is performed in accordance with a Doppler beam direction and a flow direction of the first Doppler image. Volume flow is calculated from data of the second Doppler image using angle correction determined from the first Doppler image. 
    
    
     
       In the drawings: 
         FIG.  1    illustrates the acquisition of Doppler data for volume flow assessment using pulsed Doppler with uniform insonation. 
         FIGS.  2   a  and  2   b    illustrate the acquisition of Doppler data for volume flow assessment using the Gaussian surface integration method. 
         FIG.  3    illustrates the projection of Doppler data from a Gaussian surface acquisition onto a two-dimensional image plane for quantification of volume flow. 
         FIG.  4    illustrates the acquisition of Doppler data for volume flow assessment by a transverse color Doppler method. 
         FIG.  5    illustrates the projection of biplane image planes from a 2D array transducer. 
         FIG.  6    illustrates the acquisition of Doppler data from a vessel for volume flow assessment in accordance with the principles of the present invention. 
         FIG.  7    illustrates side-by-side biplane ultrasound images during volume flow Doppler data acquisition in accordance with the present invention. 
         FIG.  7   a    illustrates the transverse view image of  FIG.  7    with the vessel lumen segmented by a circle template. 
         FIG.  8    illustrates in block diagram form an ultrasonic diagnostic imaging system constructed in accordance with the principles of the present invention. 
         FIG.  8   a    illustrates the detailed operation of the volume flow calculator of  FIG.  8   . 
         FIG.  9    is a flowchart of a method for measuring volume flow in accordance with the principles of the present invention. 
     
    
    
     Referring first to  FIG.  1   , an ultrasound image is shown depicting the acquisition of Doppler data for volume flow assessment using pulsed Doppler with uniform insonation. This is probably the most widely accepted method in clinical use, because it is available on most commercial ultrasound systems and only requires a one-dimensional (1D) array transducer. The probe operates to alternately acquire B mode echoes for production of a structural image of the tissue and blood vessels, and Doppler echoes for spatial flow velocity depiction inside a color box. The method uses pulsed Doppler ultrasound, whereby a long Doppler sample is placed at an angle to the vessel of interest and volume flow is calculated from the time-average mean velocity of the blood flow.  FIG.  1    illustrates an ultrasound image of the method, which shows a target blood vessel  70  in which volume flow is to be measured. Doppler acquisition is performed in a color box  80  which is seen to be tilted from upper left to the low right, the direction of Doppler beam transmission, and a Doppler gate line  82  is aligned with this beam direction. The color box is Doppler-scanned at the angle shown by parallel scanlines, a technique known as steered linear scanning. The distance of the break in the Doppler gate line  82 , with the vessel  70  spanned by the break, establishes the long Doppler sample across the lumen of the vessel. An adjustable flow cursor  84  is positioned over the vessel, with the top and bottom of the cursor positioned at the wall of the vessel and the intermediate horizontal lines aligned with the direction of flow. The angle between the Doppler gate line  82  and the horizontal lines of the flow cursor  84  is the angle used for Doppler angle correction. 
     However, while still widely used clinically, this method is known to be imprecise and inaccurate due to several incorrect assumptions and measurement dependencies. See, e.g., R. W. Gill,  Measurement of blood flow by ultrasound: accuracy and sources of error , Ultrasound in Medicine and Biology, vol. 11 (4), at pp 625-641 (1985). One assumption implicit in the method is that the vessel is uniformly insonated by the ultrasound beam. Since the ultrasound beam is generally smaller than the vessel in elevation, this assumption is typically not valid. If uniform insonation cannot be assumed, then a simplifying assumption must be made that the vessel cross-section is circular. This is typically true only for large arteries and usually is not true for veins. Another implicit assumption is that the temporal sampling rate of the flow is fast enough to capture the variation of flow velocity (and hence volume) through the cardiac cycle. For the pulsed Doppler method, this assumption is usually valid since the temporal sampling rate of a 1D array probe is typically adequate even for very pulsatile flow. 
     In addition, the accuracy of the measurement is very dependent on accurately determining the Doppler angle and the vessel diameter. The accuracy of the vessel diameter is important, as the diameter is used to determine the vessel cross-sectional area by calculating the vessel cross-sectional area by the equation Area=πr 2  and then multiplying the area by the flow velocity as corrected by the Doppler angle to estimate volume flow. Accurately determining the Doppler angle is relatively easy for a straight, superficial vessel, but more difficult for bending or deeper vessels. The volume measurement is particularly sensitive to the vessel diameter measurement since the diameter is used to determine the cross-sectional area by means of the above square law. 
     Other methods for assessing volume flow have been proposed that have less dependency on these assumptions and measurements. One such method is the Gaussian surface integration method, which uses 3D/4D color Doppler and Gauss&#39;s law. With this method, the flow volume is determined by integrating (summing) all the color flow voxels over a coronal surface that intersects the target vessel and is perpendicular to the 3D (or 4D) color Doppler beams. See O. D. Kripfgans et al.,  Measurement of volumetric flow , J. Ultrasound Med. vol. 25, at pp 1305-1311 (2006). See also U.S. Pat. No. 6,780,155 (Li). Since the coronal plane intersects the whole vessel, there is no assumption of uniform insonation and also no assumption about the vessel being circular. Also, neither the Doppler angle or the vessel diameter need to be measured, since the ultrasound beams transmitted are perpendicular to each point on the surface.  FIG.  2   a    illustrates a vessel  70  which is intersected by a Gaussian surface  50 . The thin plane  52  of the surface is scanned by the transmission of Doppler beams from a 2D array transducer  54 , which electronically steers the beams over the surface  50  where it intersects the vessel  70 . A flow image  76  is thereby rendered as a cross-sectional surface of the curved cross-section  58  through the vessel  70 . As explained in the foregoing &#39;155 patent, the flow image  76  can be projected onto a planar surface  72  as a B mode image  56 . The vessel lumen  62  in the B mode image can be segmented by a circle  64  or other shape outside the vessel wall  60 , and color voxels within the segmented area are then summed to produce an estimate of volume flow. Corrections should be made for color voxels at the vessel walls which may only have partial flow, and one way to do this is through normalization of the velocity estimates using the power (intensity) in the Doppler signal (see Kripfgans et al., above), often known as partial volume correction. 
     While this is an excellent method for measuring volume flow, there are challenges in measuring pulsatile flow due to the typically limited volume rates possible with 3D/4D color Doppler, resulting in under-sampled temporal information and erroneous flow volume calculation. Each point on the Gaussian surface must be sampled by an individual Doppler beam, and multiple times by multiple transmissions so as to estimate the Doppler velocity at each point on the Gaussian surface accurately. To mitigate this limitation related methods have been developed to acquire information over multiple cardiac cycles, and then either averaged to get average volume flow or, if the cardiac period is precisely known, it is possible to reconstruct a single cardiac cycle from the multiple cycles. However, these approaches add to the acquisition time and make the method less robust due to the temporal sampling time required. 
     Another method with similarities to the Gaussian surface integration method has been proposed by Picot et al. See Picot et al.,  Rapid volume flow rate estimation using transverse colour Doppler imaging , Ultrasound in Medicine and Biology, vol. 21 (9), at pp 1199-1209 (1995). In this method, instead of extracting a coronal plane from a 3D color Doppler volume, a conventional 1D array transducer is angled toward a vessel so that its scan plane, and 2D color image, intersects the vessel at an oblique but transverse angle.  FIG.  4    illustrates an ultrasound image of a vessel  70  scanned in this manner within a color box  80 . Similar to the Gaussian surface integration method, all the color Doppler pixels in the color box  80  are summed, with corrections applied for partially filled pixels at the edges of the vessel. Again, there is no assumption about flow profile or vessel geometry. One major benefit compared to the Gaussian surface integration method is that two-dimensional color frame rates are typically much higher than 3D/4D color volume rates, so adequate temporal sampling for pulsatile flow is much improved. However, one disadvantage compared to the Gaussian surface integration method is that the Doppler angle must be known, which is very difficult to obtain from a transverse image. Picot et al. describe an elaborate probe holder that allows the same vessel to be interrogated from two angles, allowing the Doppler angle dependency to be eliminated, but such a probe holder is cumbersome and impractical for clinical use. 
     In accordance with the principles of the present invention, an ultrasound probe with a two dimensional matrix array transducer is operated in the biplane mode to measure volume flow. In the biplane mode, two image planes are scanned simultaneously in an interleaved manner. While the biplane mode can be performed by a mechanical probe which moves a 1D transducer array to scan two image planes of a volumetric region as described in U.S. Pat. No. 6,443,896 (Detmer), it is preferable to use a 2D matrix array probe by which the planes are scanned electronically, rather than mechanically, as described in U.S. Pat. No. 6,709,394 (Frisa et al.) Furthermore, it is possible to perform colorflow imaging in the biplane mode as described in U.S. Pat. No. 7,645,237 (Frisa et al.) whereby a color box is scanned to acquire color Doppler data in each of the biplane image planes. Ultrasound systems and probes are commercially available which can perform colorflow scanning of biplane images, such as the xMATRIX family of probes available on Philips Healthcare ultrasound systems. In one implementation of the present invention the biplane mode of scanning is performed to generate two real-time images, including color Doppler data, that are perpendicular to each other. This allows the production, simultaneously, of a long axis view of a vessel and a transverse view. The long axis image can be used to accurately measure the Doppler angle, as described below. The transverse color image intersects the vessel obliquely, in the same manner as in Picot&#39;s method, so the same algorithm can be used to estimate the volume flow by summing all the color pixels, but with an accurately known Doppler angle correction obtained from the long axis image. 
     An implementation of the present invention overcomes many of the limitations and shortcomings of the prior methods for measuring volume flow. As compared to the pulsed Doppler method, an implementation of the present invention requires no assumptions of uniform insonation or vessel geometry, and there is no need to measure the vessel diameter, which is the greatest cause of inaccuracy in the typical pulsed Doppler-based method. As compared to the Gaussian surface integration method, the inventive technique provides much better temporal sampling since only two image planes need be scanned and so is more suitable for very pulsatile flow as would be found in many arteries. In addition, spatial sampling is uncompromised, as there is no need to try to improve 3D volume frame rates. Better spatial sampling results in better representation of the flow profile and also reduced reliance on partial volume correction. As compared to the method of Picot et al., an implementation of the present invention makes it very easy to accurately measure the Doppler angle needed for velocity correction. 
     The operation and use of an implementation of the present invention may be appreciated by referring to  FIGS.  5  and  6   .  FIG.  5    depicts a 2D matrix array transducer  54  which scans two biplanes in front of the transducer, denoted L and T. When all of the scanlines transmitted and received for a plane are normal to the plane of the 2D array transducer, the plane will extend normal to the transducer as shown in this illustration. When the scanlines are transmitted and received at an oblique angle to the plane of the array transducer, the scan plane will be angled in the shape of a parallelogram, which results from steered linear operation. In the example of  FIG.  5    the two biplanes extend normal to the transducer, as the scanlines are transmitted and received straight ahead of the array. The L and T planes are seen to intersect at a common intermediate scanline. 
       FIG.  6    shows the two L and T biplanes being steered to intersect and scan a blood vessel  70  in accordance with the present invention. The L plane is tilted from the upper left to the lower right and is aimed by probe manipulation at the longitudinal center of the vessel  70 , thereby producing a long axis view of the vessel. The L scan plane is tilted in a parallelogram orientation so that the scanlines of the plane will intersect the direction of blood flow of the vessel at a nonorthogonal angle since, as is well known, a 90° angle between a Doppler beam and the flow direction will yield no measurable Doppler signal since the cosine of 90° is zero. The T plane is aligned with one of the parallel scanlines of the L plane and intersects the blood vessel  70  obliquely, thereby producing a transverse view of the vessel where the T image plane cuts through the vessel. The Doppler angle of the T plane which is needed for angle correction of the Doppler velocity measurements is thus the angle at which the L plane is tilted, which is readily known from the Doppler angle of the L plane and the readily observable flow direction in the long axis view of the vessel. 
     When a biplane probe is used to scan a vessel as illustrated in  FIG.  6   , long axis and transverse views of the vessel can be produced and displayed simultaneously as illustrated in  FIG.  7   . In this example the biplane images are shown side-by-side in a duplex display. The left image  90  is the long axis view of the L scan plane of  FIG.  6   , showing vessel  70  in a long axis view bisecting the vessel. Within the grayscale (B mode) image of the blood vessel and surrounding tissue is a color box  80  which is scanned by Doppler beams for Doppler display of the material inside the box. Like the color box of  FIG.  1   , the color box  80  is tilted at an angle as established by the setting of the tilt angle of the Doppler gate line  82 . Like the previous drawing, the Doppler line has a flow cursor  84 , which the user aligns with the direction of the flow in vessel  70 . The angle between the Doppler gate line  82  and the flow cursor  84  establishes the Doppler angle, the angle between the Doppler beams used to scan the color box  80  and the direction of the blood flow. The Doppler angle is commonly recognized and recorded automatically in standard ultrasound systems. 
     In an implementation of the present invention the image  92  of the transverse view of the vessel  90  is scanned in a plane in alignment with the angle of the color box  80 . Typically, the image planes  90  and  92  are spatially normal to each other. In this example the plane of the image  92  is in alignment with the Doppler gate line  82  of image  90 ; the two images spatially share a common location of their Doppler lines  82 . The result is that the Doppler angle correction needed for the velocity values of flow in the transverse image  92  is the Doppler angle of the longitudinal view  90 , the angle between the Doppler gate line  82  and the flow cursor  84 , which is readily recognized in a typical commercial ultrasound system. The ultrasound system can then measure volume flow by any of several known algorithms such as that of Picot et al., in which the color pixel values of the vessel in the transverse view are angle-corrected, then summed to compute volume flow. Mathematically, this can be represented by Gauss&#39;s theorem, calculated as: 
     
       
      
       Q=∫ 
       S 
       v·dA  
      
     
     where Q is the volume flow in, e.g., milliliters per second, v is angle-corrected flow velocity, and the surface S is the Doppler portion of the cut plane through the vessel lumen in the transverse view  92 . In addition, a typical commercial ultrasound system will enable a user to segment (delineate) the portion of the image over which Doppler velocity pixels are to be integrated.  FIG.  7   a    presents an example of such a tool, a circle template  78  which a user can appropriately size and then maneuver over an ultrasound image to designate the image area within which Doppler value integration is to occur. Such segmentation of the vessel lumen will prevent the volume flow algorithm from mistakenly including pixel values of neighboring blood vessels, for instance. 
     In  FIG.  8   , an ultrasound system constructed in accordance with the principles of the present invention is shown in block diagram form. A transducer array  12  is provided in an ultrasound probe  10  for transmitting ultrasonic waves and receiving echo information. The transducer array  12  is a two-dimensional array of transducer elements capable of scanning in three dimensions, in both elevation and azimuth. It is thus capable of scanning two biplanes simultaneously in a time-interleaved manner. The transducer array  12  is coupled to a microbeamformer  14  in the probe which controls transmission and reception of signals by the array elements. Microbeamformers are capable of at least partial beamforming of the signals received by groups or “patches” of transducer elements as described in U.S. Pat. No. 5,997,479 (Savord et al.), U.S. Pat. No. 6,013,032 (Savord), and U.S. Pat. No. 6,623,432 (Powers et al.) The microbeamformer is coupled by the probe cable to a transmit/receive (T/R) switch  16  which switches between transmission and reception and protects the main beamformer  18  from high energy transmit signals. The transmission of ultrasonic beams from the transducer array  12  under control of the microbeamformer  14  is directed by a beamformer controller  17  coupled to the T/R switch and the main beamformer  18 , which receives input from the user&#39;s operation of the user interface or control panel  38 . Among the transmit characteristics controlled by the transmit controller are the direction, number, spacing, amplitude, phase, angle, frequency, polarity, and diversity of transmit waveforms. Beams formed in the direction of pulse transmission may be steered straight ahead from the transducer array, or at different angles on either side of an unsteered beam for a wider sector field of view, or for transmission at a selected Doppler angle. 
     The echoes received by a contiguous group of transducer elements are beamformed by appropriately delaying them and then combining them. The partially beamformed signals produced by the microbeamformer  14  from each patch are coupled to the main beamformer  18  where partially beamformed signals from individual patches of transducer elements are combined into a fully beamformed coherent echo signal. For example, the main beamformer  18  may have 128 channels, each of which receives a partially beamformed signal from a patch of 12 transducer elements. In this way the signals received by over 1500 transducer elements of a two-dimensional matrix array transducer can contribute efficiently to a single beamformed signal. 
     The coherent echo signals undergo signal processing by a signal processor  20 , which includes filtering by a digital filter and noise and speckle reduction as by spatial or frequency compounding. The digital filter of the signal processor  20  can be a filter of the type disclosed in U.S. Pat. No. 5,833,613 (Averkiou et al.), for example. The echo signals are then coupled to a quadrature bandpass filter (QBP)  22 . The QBP performs three functions: band limiting the r.f. echo signal data, producing in-phase and quadrature pairs (I and Q) of echo signal data, and decimating the digital sample rate. The QBP comprises two separate filters, one producing in-phase samples and the other producing quadrature samples, with each filter being formed by a plurality of multiplier-accumulators (MACs) implementing an FIR filter. 
     The beamformed and processed coherent echo signals are coupled to a pair of image data processors. A B mode processor  26  produces signal data for a B mode image of structure in the body such as tissue and blood vessel walls. The B mode processor performs amplitude (envelope) detection of quadrature demodulated I and Q signal components by calculating the echo signal amplitude in the form of (I 2 +Q 2 ) 1/2 . The quadrature echo signal components are also coupled to a Doppler processor  24 . The Doppler processor  24  stores ensembles of echo signals from discrete points in an image field which are then used to estimate the Doppler shift at points in the image with a fast Fourier transform (FFT) processor. The Doppler processor can also perform angle correction of Doppler velocity values, and in an implementation of the present invention angle correction as measured on a first (long axis) Doppler image is used to perform angle correction of the Doppler data of a second (transverse) Doppler image used to determine volume flow. The rate at which the ensembles are acquired determines the velocity range of motion that the system can accurately measure and depict in an image. The Doppler shift is proportional to motion at points in the image field, e.g., blood flow and tissue motion. For color Doppler image data, the estimated Doppler flow values at each point in a blood vessel are wall filtered, angle corrected, and converted to color values using a look-up table. The wall filter has an adjustable cutoff frequency above or below which motion will be rejected such as the low frequency motion of the wall of a blood vessel when imaging flowing blood. The B mode image data and the Doppler flow values are coupled to a scan converter  28  which converts the B mode and Doppler samples from their acquired R-θ coordinates to Cartesian (x,y) coordinates for display in a desired display format, e.g., a rectilinear display format or a sector display format as shown in  FIGS.  7  and  7     a . Either the B mode image or the Doppler image may be displayed alone, or the two shown together in anatomical registration in which the color Doppler overlay shows the blood flow in B mode processed tissue and vessels in the image as shown in  FIGS.  7  and  7     a . Another display possibility is to display side-by-side images of the same anatomy which have been processed differently, as shown in these drawings. This side-by-side display format is useful when comparing images, and is particularly useful for displaying both images of a biplane probe. The scan-converted image data, both B mode and Doppler data, is coupled to and stored in an image data memory  30  where it is stored in memory locations addressable in accordance with the spatial locations from which the image data values were acquired. The biplane images from data produced by the scan converter  28  and stored in the image data memory are coupled to a display processor  34  for further enhancement, buffering and temporary storage for display on an image display  36 . 
     The Doppler values of an image, such as the color Doppler pixel values of a transverse view of a blood vessel as shown in image  92  of  FIG.  7   , are coupled to a volume flow calculator  40 . There, an algorithm computing volume flow is executed, such as the integration of pixel flow velocity values over the area of a vessel lumen as exemplified by the expression 
     
       
      
       Q=∫ 
       S 
       v·dA  
      
     
     Volume flow may be calculated for each frame separately in the color Doppler image to provide volume flow as a function of time, or may be summed over multiple frames to provide a time-average volume flow rate. The volume flow measurement is coupled to a graphics generator  49 , from which the volume flow value is coupled to the display processor  34  for display in conjunction with the ultrasound images. Alternatively or additionally, the graphics generator  49  can produces a flow profile curve for display on the display  36 . The graphics generator also produces graphics for display with the ultrasound image for things such as cursors, measurement dimensions, exam parameters, patient name, and the aforementioned Doppler gate line  82 , flow cursor  84 , and segmentation template  78 . 
     Details of the operation of the volume flow calculator  40  of  FIG.  8    are illustrated in  FIG.  8   a   . In this implementation, angle correction of the transverse image  92  is performed by the volume flow calculator rather than the Doppler processor  24 . An image segmentation processor  402  receives the data of transverse image  92  from the image data memory  30 . The Doppler data in the transverse view of vessel  70  is identified (segmented) by the placement by the system operator of a template such as a circle template  78  over the lumen of vessel  70 . The Doppler data thereby designated within the vessel  70  is coupled to an angle correction processor  404 , which computes the angle correction performed for Doppler data of the long axis view  90  from the long axis Doppler gate line  82  and the long axis flow cursor  84 , which is the angle between those two graphics as set by the user during acquisition of the long axis view  90 . The Doppler values of the flow in the transverse image are thereby angle corrected in accordance with the angle set by these graphics during acquisition of the long axis view  90 . The angle-corrected Doppler values are then integrated over the vessel area in the transverse image  92  to determine the volume flow. In the implementation of  FIG.  8   a    this operation is performed by summing the angle-corrected Doppler flow values of the transverse view image as illustrated by processor  406  to produce a measure Q of volume flow. This value is coupled to the graphics generator  49  for display to the ultrasound system operator. In a typical implementation the computation performed by the processors  402 ,  404 , and  406  are performed by software programs configured to perform the illustrated functions. 
     The ultrasound system of  FIG.  8    can be operated to perform a typical ultrasound carotid artery exam as follows. The user would first examine the carotid artery by ultrasound imaging, including B-mode, color Doppler and pulsed Doppler imaging in accordance with standard clinical practice. On identifying a significant stenosis, the user then takes the following steps to determine the volume flow. First, scan with the ultrasound probe to find a section of the artery upstream of the stenosis to minimize turbulence in the blood flow. Next, biplane mode is actuated. The probe is manipulated until the vessel appears stretched out in a long axis view on the left side image  90 . The probe is tilted in elevation until the vessel is roughly in the center of the right side image  92 . The color scale (pulse repetition frequency, PRF) is adjusted so that the flow is not aliased. The angle correction is adjusted so that it aligns with the vessel  70  in the left side image  90 . One or more cardiac cycles of flow information are then acquired. A region of interest is designated (segmented) by placing a template  78  over the vessel  70  in the right side image to define which color pixels in the image should be included in the volume flow calculation. Volume flow is then calculated and displayed using Doppler data of the transverse image as angle-corrected from the angle correction of the long axis view, either as an average over time (a single number) or as a graph of volume flow as a function of time. 
     A method for measuring volume flow in accordance with the principles of the present invention may be conducted as shown in  FIG.  9   . At the start  901 , an ultrasound system is set to operate in the biplane mode, acquiring two images which intersect at a Doppler angle of one of them. At  903 , a first Doppler image is acquired such as the long axis view  90  of  FIG.  7   . The direction of Doppler acquisition of this first image is adjusted at  904  by adjusting a Doppler gate line  82 . After the Doppler direction has been set, a second Doppler image is acquired at  905  at a plane in the Doppler line direction of the first image. At  907  both images are simultaneously displayed. The Doppler angle correction for the first image is determined at  909  from the Doppler line direction and the direction of a flow cursor set by the user over the flow of the first image. At  911  the flow of the second Doppler image is segmented as by placing a graphic circle template over the lumen of the second image. At  915  volume flow is calculated as described above from the Doppler values of the flow of the second image as angle-corrected by the angle correction determine for the first image. At  920  the measured volume flow is displayed to the user or recorded in the ultrasound exam record. 
     It is noted that the scope of the invention described above also includes embodiments which do not necessarily include an ultrasound probe, but which instead receives an input of acquired Doppler image data from two image planes ( 90 ,  92 ) intersecting along a Doppler beam direction and adapted to produce two Doppler images of flow. The invention further includes a display ( 36 ) adapted to display the two Doppler images simultaneously, and a graphics generator ( 49 ), responsive to a user control, and adapted to display a Doppler line ( 82 ) and a flow cursor ( 84 ) over a first one ( 90 ) of the Doppler images. A volume flow calculator ( 40 ) is responsive to Doppler image data of the second one ( 92 ) of the Doppler images and a Doppler angle established by the Doppler line and the flow cursor, which is adapted to determine an angle-corrected measure of volume flow. 
     It should further be noted that an ultrasound system suitable for use in an implementation of the present invention, and in particular the component structure of the ultrasound system of  FIG.  8   , may be implemented in hardware, software or a combination thereof. The various embodiments and/or components of an ultrasound system and its controller, or components and controllers therein, also may be implemented as part of one or more computers or microprocessors. The computer or processor may include a computing device, an input device, a display unit and an interface, for example, for accessing the internet. The computer or processor may include a microprocessor. The microprocessor may be connected to a communication bus, for example, to access a PACS system or the data network for importing training images and storing the results of clinical exams. The computer or processor may also include a memory. The memory devices such as the image data memory may include Random Access Memory (RAM) and Read Only Memory (ROM). The computer or processor further may include a storage device, which may be a hard disk drive or a removable storage drive such as a floppy disk drive, optical disk drive, solid-state thumb drive, and the like. The storage device may also be other similar means for loading computer programs or instructions for volume flow analysis into the computer or processor. 
     As used herein, the term “computer” or “module” or “processor” or “workstation” may include any processor-based or microprocessor-based system including systems using microcontrollers, reduced instruction set computers (RISC), ASICs, logic circuits, and any other circuit or processor capable of executing the functions described herein. The above examples are exemplary only and are thus not intended to limit in any way the definition and/or meaning of these terms. 
     The computer or processor executes a set of instructions that are stored in one or more storage elements, in order to process input data. The storage elements may also store data or other information as desired or needed. The storage element may be in the form of an information source or a physical memory element within a processing machine. The set of instructions of an ultrasound system including those controlling the acquisition, processing, and display of ultrasound images and instructions for Doppler angle measurement and volume flow calculation as described above may include various commands that instruct a computer or processor as a processing machine to perform specific operations such as the methods and processes of Doppler flow data acquisition, line and cursor adjustment, and volume flow measurement. The set of instructions may be in the form of a software program. The software may be in various forms such as system software or application software and which may be embodied as a tangible and non-transitory computer readable medium. The equation given above for volume flow calculation and the summation of Doppler data values shown in  FIG.  8   a   , as well as the calculation of the Doppler angle from the cursors placed over an image, are typically calculated by or under the direction of software routines. Further, the software may be in the form of a collection of separate programs or modules within a larger program or a portion of a program module. The software also may include modular programming in the form of object-oriented programming. The processing of input data by the processing machine may be in response to operator commands issued from control panel  38 , or in response to results of previous processing, or in response to a request made by another processing machine. 
     Furthermore, the limitations of the following claims are not written in means-plus-function format and are not intended to be interpreted based on 35 U.S.C. 112, sixth paragraph, unless and until such claim limitations expressly use the phrase “means for” followed by a statement of function devoid of further structure.