WO2018058577A1 - 超声血流的参数显示方法及其超声成像系统 - Google Patents

超声血流的参数显示方法及其超声成像系统 Download PDF

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Publication number
WO2018058577A1
WO2018058577A1 PCT/CN2016/101208 CN2016101208W WO2018058577A1 WO 2018058577 A1 WO2018058577 A1 WO 2018058577A1 CN 2016101208 W CN2016101208 W CN 2016101208W WO 2018058577 A1 WO2018058577 A1 WO 2018058577A1
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Prior art keywords
blood flow
flow velocity
ultrasonic
dispersion
ultrasound
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PCT/CN2016/101208
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English (en)
French (fr)
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杜宜纲
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深圳迈瑞生物医疗电子股份有限公司
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Priority to CN201680084275.5A priority Critical patent/CN108882916B/zh
Priority to CN202210535563.5A priority patent/CN114848016A/zh
Priority to PCT/CN2016/101208 priority patent/WO2018058577A1/zh
Publication of WO2018058577A1 publication Critical patent/WO2018058577A1/zh
Priority to US16/367,514 priority patent/US11304677B2/en
Priority to US17/723,107 priority patent/US11890141B2/en

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    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
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    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/06—Measuring blood flow
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/13—Tomography
    • A61B8/14—Echo-tomography
    • A61B8/145—Echo-tomography characterised by scanning multiple planes
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    • A61B8/44—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device
    • A61B8/4477—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device using several separate ultrasound transducers or probes
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    • A61B8/4483—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device characterised by features of the ultrasound transducer
    • A61B8/4494—Constructional features of the ultrasonic, sonic or infrasonic diagnostic device characterised by features of the ultrasound transducer characterised by the arrangement of the transducer elements
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Definitions

  • the invention relates to a blood flow information imaging display technology in an ultrasound system, in particular to a parameter display method of an ultrasonic blood flow and an ultrasonic imaging system thereof.
  • Color Doppler flowmetry is the same as pulse wave and continuous wave Doppler, and is also imaged by Doppler effect between red blood cells and ultrasound.
  • Color Doppler flowmeter includes two-dimensional ultrasound imaging system, pulse Doppler (one-dimensional Doppler) blood flow analysis system, continuous wave Doppler blood flow measurement system and color Doppler (two-dimensional Doppler) Blood flow imaging system.
  • the oscillator generates two orthogonal signals with a phase difference of ⁇ /2, which are respectively multiplied by the Doppler blood flow signal, and the product is converted into a digital signal by an analog/digital (A/D) converter, and filtered by a comb filter.
  • A/D analog/digital
  • the autocorrelator After removing the low frequency component generated by the blood vessel wall or the valve, it is sent to the autocorrelator for autocorrelation detection. Since each sample contains Doppler blood flow information generated by many red blood cells, a mixed signal of multiple blood flow velocity vectors is obtained after autocorrelation detection.
  • the autocorrelation test result is sent to the speed calculator and the variance calculator to obtain an average speed, and is stored in the digital scan converter (DSC) together with the FFT-processed blood flow spectrum information and the two-dimensional image information.
  • DSC digital scan converter
  • the blood flow data is encoded as a pseudo color by the color processor, and sent to the color display for color Doppler blood flow imaging.
  • Spectral Doppler is used for quantitative diagnosis of heart valve stenosis and arteriosclerotic lesions.
  • the blood flow may be different at different times in a cardiac cycle.
  • the blood flow of the common carotid artery is laminar under normal conditions, but if plaque occurs and the arterial stenosis occurs, the blood flow becomes more disordered.
  • Eddy currents may occur near the stenosis during systole.
  • the degree of eddy current is also an important indicator for judging the stenosis rate.
  • the eddy current area is used as the judgment index of the eddy current degree.
  • the traditional color Doppler cannot measure the direction of blood flow, and can only be manually depicted by red and blue colors and related clinical experience. The area of the eddy current is out, so that errors are easily generated.
  • the degree of eddy current needs to be quantitatively calculated to make the diagnosis more reliable.
  • a parameter display method for ultrasonic blood flow which includes:
  • the quantized result of the dispersion is displayed.
  • an ultrasound imaging system comprising:
  • a receiving circuit and a beam combining module configured to receive an echo signal of the ultrasonic beam, and perform beam combining to obtain an ultrasonic signal
  • An image processing module configured to obtain a blood flow velocity direction in the scan target according to the ultrasonic signal, extract a plurality of blood flow velocity directions, and quantize the dispersion of the extracted plurality of blood flow velocity directions;
  • a display for displaying the quantized result of the dispersion.
  • FIG. 1 is a block diagram showing an ultrasonic imaging system according to an embodiment of the present invention
  • FIG. 2 is a schematic diagram of a vertically emitted planar ultrasonic beam according to an embodiment of the present invention
  • FIG. 3 is a schematic diagram of a deflected-emitting planar ultrasonic beam according to an embodiment of the present invention
  • FIG. 4 is a schematic diagram of multi-angle reception in an embodiment of the present invention.
  • FIG. 5 is a schematic flowchart of a method according to an embodiment of the present invention.
  • FIG. 6 is a schematic diagram of a calculation method of a blood flow velocity vector according to an embodiment of the present invention.
  • Figure 7 is a schematic diagram showing the spatial coordinate representation of the direction of blood flow velocity in an embodiment of the present invention.
  • FIG. 8 is a schematic diagram of calculation of a dispersion image in one embodiment of the present invention.
  • 9(a) is a schematic diagram of calculation of blood flow velocity vector information in a first mode in one embodiment of the present invention.
  • Figure 9 (b) is a schematic diagram of calculation of blood flow velocity vector information in the second mode in one embodiment of the present invention.
  • FIG. 10 is a schematic flowchart of a method according to an embodiment of the present invention.
  • FIG. 11 is a schematic flow chart of a method according to an embodiment of the present invention.
  • FIG. 12 is a schematic flow chart of a method according to an embodiment of the present invention.
  • FIG. 13 is a schematic diagram showing a display effect of a dispersion image according to an embodiment of the present invention.
  • Figure 14 is a schematic view of the line of Figure 13;
  • Figure 16, Figure 17, Figure 18, and Figure 19 are schematic illustrations of comparative displays of ultrasound images and discrete quantized structures, respectively, in various embodiments of the present invention.
  • the ultrasonic imaging system generally includes a probe 1, a transmitting circuit 2, a transmitting/receiving selection switch 3, a receiving circuit 4, a beam combining module 5, a signal processing module 6, an image processing module 7, and a display 8.
  • “Multiple” in this document means 2 or more.
  • the transmitting circuit 2 transmits a delayed-focused transmission pulse having a certain amplitude and polarity to the probe 1 through the transmission/reception selection switch 3.
  • the probe 1 is excited by a transmitting pulse to transmit an ultrasonic wave to a scanning target (for example, an organ, a tissue, a blood vessel, or the like in a human body or an animal body, not shown), and receives a reflection from the target area after a certain delay.
  • the ultrasonic echo signal of the target information is scanned and the ultrasonic echo signal is reconverted into an electrical signal.
  • the receiving circuit receives the electrical signals generated by the conversion of the probe 1 and sends the ultrasonic echo signals to the beam combining module 5.
  • the beam synthesizing module 5 performs processing such as focus delay, weighting, and channel summation on the ultrasonic signal to obtain an ultrasonic signal, and then sends the ultrasonic signal to the signal processing module 6 for related signal processing, such as filtering.
  • the ultrasonic signal processed by the signal processing module 6 is sent to the image processing module 7.
  • the image processing module 7 performs different processing on the signals according to different imaging modes required by the user, obtains image data of different modes, and then forms ultrasonic images of different modes by logarithmic compression, dynamic range adjustment, digital scan conversion, etc., such as A two-dimensional image such as a B image, a C image, or a D image, and in addition, the ultrasonic image may further include a three-dimensional image.
  • the ultrasonic image generated by the image processing module 7 is sent to the display 8 for display.
  • the image processing module 7 may further calculate a blood flow velocity vector of the target point within the scan target according to the ultrasonic signal, and may add the calculated blood flow velocity vector rendering to the display for display by the rendering process, and/or The calculated blood flow velocity vector is sent to the display for display of speed information.
  • the image processing module 7 and the signal processing module 6 are provided separately on different processors or integrated on the same processor 9.
  • the target point mentioned in this embodiment may be one pixel point on the ultrasound image or a region block containing at least two pixel points.
  • the blood flow velocity vector of the target point including the velocity value and the velocity direction, is used to characterize the flow velocity information of the blood flow motion state within the scanning target. The relevant calculation of the blood flow velocity vector will be explained in detail below.
  • Probe 1 typically includes an array of multiple array elements.
  • Each of the array elements of the probe 1 or a portion of all of the array elements participate in the transmission of the ultrasonic waves each time an ultrasonic wave is transmitted or an ultrasonic wave is received.
  • each of the array elements or each of the array elements participating in the ultrasonic transmission are respectively excited by the transmitting pulse, and respectively emit ultrasonic waves, and the ultrasonic waves respectively emitted by the array elements are superimposed during the propagation to form the emitted
  • the propagation direction of the synthetic ultrasonic beam is the emission angle of the ultrasonic wave mentioned herein.
  • the array elements participating in the ultrasonic transmission may be excited by the transmitting pulse at the same time; or, there may be a certain delay between the time when the array elements participating in the ultrasonic transmission are excited by the transmitting pulse. Participation in ultrasound by control
  • the delay between the time at which the transmitted elements of the wave are excited by the transmitted pulse can change the direction of propagation of the synthesized ultrasonic beam (i.e., the angle of emission), as will be described in more detail below.
  • the ultrasonic waves emitted by the respective array elements participating in the transmission of the ultrasonic waves will not be focused during the propagation, nor will they completely diverge. It is a plane wave that is generally planar as a whole.
  • the ultrasonic beams emitted by the respective array elements can be superimposed at predetermined positions, so that the intensity of the ultrasonic waves is maximum at the predetermined position, that is, The ultrasonic waves emitted by the respective array elements are "focused" to the predetermined position, the predetermined position of the focus being referred to as the "focus", such that the resulting synthesized ultrasonic beam is a beam focused at the focus, referred to herein as " Focus on the ultrasound beam.”
  • the array elements participating in the transmission of the ultrasonic waves may operate in a manner that has a predetermined transmission delay (ie, there is a predetermined delay between the time when the array elements participating in the transmission of the ultrasonic waves are excited by the emission pulse), each The ultrasonic waves emitted by the array elements are focused at the focus to form a focused ultrasound beam.
  • the ultrasonic waves emitted by the respective array elements participating in the emission of the ultrasonic waves are diverged during the propagation, forming a substantially divergent overall. wave.
  • the ultrasonic wave of this divergent form is referred to as a "divergent ultrasonic beam.”
  • a plurality of array elements arranged linearly are simultaneously excited by an electric pulse signal, and each array element simultaneously emits ultrasonic waves, and the propagation direction of the synthesized ultrasonic beam is consistent with the normal direction of the array plane of the array elements.
  • the plane wave emitted vertically there is no time delay between the array elements participating in the transmission of the ultrasonic wave (that is, there is no delay between the time when each array element is excited by the transmitting pulse), and each array element is emitted with a pulse.
  • the generated ultrasonic beam is a plane wave, that is, a plane ultrasonic beam, and the propagation direction of the plane ultrasonic beam is substantially perpendicular to the surface of the probe 1 from which the ultrasonic wave is emitted, that is, the propagation direction of the synthesized ultrasonic beam and the normal direction of the arrangement plane of the array element The angle between them is zero degrees.
  • the excitation pulse applied to each array element has a time delay, and each array element sequentially emits an ultrasonic beam according to the time delay, the propagation direction of the synthesized ultrasonic beam and the normal direction of the array element arrangement plane are With a certain angle, that is, the transmission angle of the composite beam, changing the above time delay, the size of the transmission angle of the composite beam and the emission in the normal direction of the array plane of the array element can be adjusted.
  • FIG. 3 shows a plane wave that is deflected and emitted.
  • the respective array elements participating in the transmission of the ultrasonic wave that is, there is a predetermined time delay between the time when each array element is excited by the transmitting pulse
  • the array elements are excited by the transmitted pulses in a predetermined order.
  • the generated ultrasonic beam is a plane wave, that is, a plane ultrasonic beam, and the propagation direction of the plane ultrasonic beam is at an angle to the normal direction of the array arrangement plane of the probe 1 (for example, the angle a in FIG. 3), and the angle is The angle of emission of the ultrasonic beam of the plane.
  • the composite beam here may be the planar ultrasonic beam, the focused ultrasonic beam or the divergent ultrasonic beam mentioned above. .
  • the two-dimensional ultrasonic transducer can be understood as a combination of a plurality of linear arrays, and therefore, the two-dimensional ultrasonic transducer can also be excited by the emission of the array elements participating in the transmission of the ultrasonic waves.
  • the delay between the adjustments is to adjust the "emission angle" of the composite beam formed between the composite beam and the normal direction of the array of elements.
  • the ultrasonic probe includes the array unit 1, the array unit 2, the array unit 3, and the array unit 4.
  • the array unit 1, the array unit 2, the array unit 3, and the array unit 4 may be one array element or a plurality of array elements. A combination of one or more of the array unit 1, the array unit 2, the array unit 3, and the array unit 4 may be used as the receiving element.
  • the array element 1 when an ultrasonic beam of a transmission angle is transmitted to a scanning target including the target point position A, the array element 1 is used as a receiving element, and the ultrasonic beam reflected back from a certain target point position A in the scanning target is received.
  • the wave based on the line connecting the aperture position of the element unit 1 and the target point position A (marked as a solid line in Fig. 4), can determine the reception angle a1 of the echo of the ultrasonic beam received at the current time.
  • the echo of the ultrasonic beam reflected from a certain target point position A in the scanning target is received, according to the connection between the aperture position of the array unit 2 and the target point position A ( The dotted line in Fig. 4 can determine the reception angle a2 of the echo of the ultrasonic beam received at the current time.
  • the echo of the ultrasonic beam returned from the same target position A can obtain echoes of the ultrasonic beams of two different receiving angles. Therefore, the "receiving angle" of the echo of the ultrasonic beam is defined in accordance with the angle between the line connecting the aperture position of the receiving element and the position of the target point and the normal direction of the plane of the array of ultrasonic elements.
  • the receiving angle of the echo of the ultrasonic beam can be changed by changing the position of the aperture of the receiving element on the probe, thereby obtaining Ultrasonic signals corresponding to different receiving angles of the scanning target; can also change the emission angle of the ultrasonic beam by controlling the delay between the time when the array elements participating in the transmission of the ultrasonic waves are excited by the transmitting pulse, and the ultrasonic beams based on different emission angles Echo, obtaining ultrasonic signals corresponding to different emission angles of the scanning target.
  • the image processing module 7 can calculate a blood flow velocity vector of a plurality of target points in the region of interest in the scan target or in the scan target according to the ultrasonic signals of different angles.
  • the ultrasonic imaging system shown in FIG. 1 further includes an operation control module 10 for receiving from the operation.
  • the adjustment signal includes adjustment of imaging parameters such as an emission angle of the ultrasonic beam, a reception angle, an ultrasonic beam type, or the like, or may also include an image of the tissue image processing module, a region of interest, or a blood flow velocity
  • the vector's calculation results are adjusted.
  • the operation control module 10 can be a human-computer interaction interface, such as a keyboard, a scroll wheel, a touch gesture receiving and calculating module connected to a touch screen with a touch function, a mouse, a transceiver module for a gesture control signal, and the like.
  • the display 8 in FIG. 1 includes one or more display screens, and the display screen in this embodiment may be a touch display screen, an LED display screen, or the like.
  • the image data or the quantized result output by the image processing module may also be transmitted to the remote display through the wireless transmission module for display.
  • the solution of the embodiment is not limited to the desktop ultrasound device, and may also include all available for display in the medical internet system. Ultrasound image of the device.
  • FIG. 5 provides a parameter display method for ultrasonic blood flow, which provides a method for evaluating the degree of eddy current or turbulence in blood flow in a blood vessel, and can be used as a more intuitive quantitative analysis method for determining the degree of vascular stenosis. . The details will be described below.
  • step S100 in FIG. 5 the receiving ultrasonic wave signal from the scanning target is obtained by the probe 1 by the receiving circuit 4 and the beam combining module 5.
  • the probe 1 is excited by the transmitting circuit 2 to emit an ultrasonic beam to the scanning target, and the echo of the ultrasonic beam is received to obtain the ultrasonic signal mentioned in the step S100.
  • the ultrasonic beam emitted to the scanning target in this embodiment may include: a focused ultrasonic beam and a non-focused ultrasonic beam, wherein the unfocused ultrasonic beam includes: a virtual source ultrasonic beam, a non-diffracting ultrasonic beam, a divergent ultrasonic beam, or a planar ultrasonic beam; At least one of beam types or a combination of at least two or more (herein "the above” includes the number, the same applies hereinafter).
  • the ultrasonic signal in step S100 can be an echo signal of the ultrasonic beam.
  • the step S100 includes: Step 121: transmitting a focused ultrasound beam to the scanning target, receiving an echo of the focused ultrasound beam, obtaining a focused ultrasound signal, reconstructing the ultrasound image, or calculating a blood flow velocity vector, etc. .
  • step 122 is included: transmitting a planar ultrasonic beam to the scanning target, receiving an echo of the planar ultrasonic beam, obtaining a planar ultrasonic signal for reconstructing the ultrasonic image, or calculating a blood flow velocity vector or the like.
  • step 121 and step 122 are included in step S100: transmitting a focused ultrasonic beam to the scanning target to obtain a focused ultrasonic signal; and transmitting a planar ultrasonic beam to the scanning target to obtain a planar ultrasonic signal.
  • the focused ultrasound signal can be used to reconstruct at least a portion of the ultrasound image of the scan target to obtain a better quality ultrasound image as the background image, and the planar ultrasound signal can also be used as the calculation of the blood flow velocity vector in step S200 of FIG. Image data foundation.
  • two types of beam are used for transmission in step S100, two types of ultrasonic beams are alternately transmitted to the scanning target. For example, insertion into a scanning target during the process of transmitting a planar ultrasonic beam to a scanning target The process of focusing the ultrasonic beam, that is, performing the above steps 121 and 122 alternately. This can ensure the synchronization of the acquisition of the two ultrasound beam image data, and improve the accuracy of the blood flow velocity vector obtained by the multi-beam angle transmission.
  • ultrasonic signals of a plurality of angles can also be received in step S100 for calculating a blood flow velocity vector or an ultrasound image.
  • ultrasonic beams of different emission angles may be transmitted to the scanning target for receiving ultrasonic signals corresponding to the plurality of emission angles.
  • ultrasonic signals corresponding to different reception angles are received from the scanning target. It can be seen that the ultrasonic signals of multiple angles can correspond to multiple emission angles or multiple reception angles. The details are as follows:
  • ultrasonic signals of multiple angles can be received along different transmission angles.
  • the method may include: transmitting an ultrasonic beam to the scanning target along a plurality of emission angles, and receiving an echo of the ultrasonic beam to obtain an ultrasonic signal corresponding to the plurality of emission angles, which is obtained as the receiving in step S100. Ultrasonic signals at multiple angles.
  • the ultrasonic beam is transmitted to the scanning target along a plurality of emission angles, and in the process, the process of transmitting the ultrasonic beam to the scanning target may be alternately performed in accordance with the difference in the emission angle. For example, if an ultrasonic beam is emitted to the scanning target along two emission angles, the ultrasonic beam is first transmitted to the scanning target along the first emission angle, and then the ultrasonic beam is emitted to the scanning target along the second emission angle to complete a scanning cycle. Finally, the above scanning cycle process is sequentially repeated.
  • the ultrasonic beam may be first transmitted to the scanning target along one emission angle, and then the ultrasonic beam is transmitted to the scanning target along another emission angle, and the scanning process is completed after all the emission angles are sequentially executed.
  • it can be obtained by changing the delay time of each array element or each partial array element in the array elements participating in the ultrasonic transmission, and can be specifically explained with reference to FIG. 2 or FIG.
  • a plurality of ultrasonic beams are transmitted to the scanning target at each of the emission angles to obtain a plurality of ultrasonic signals for subsequent processing of the ultrasonic image data.
  • a plurality of unfocused ultrasonic beams are respectively emitted to the scanning target at a plurality of emission angles, or a plurality of focused ultrasonic beams are respectively transmitted to the scanning target along a plurality of emission angles. And each time the ultrasonic beam is emitted, an ultrasonic signal is obtained.
  • the process of transmitting a plurality of ultrasonic beams to the scanning target is alternately performed according to the difference of the emission angles, so that the obtained echo data can approximate the blood flow velocity vector of the target point at the same time, and the calculation accuracy of the velocity vector information is improved. For example, if N times of ultrasonic beams are respectively transmitted to the scanning target along three emission angles, at least one ultrasonic beam may be transmitted to the scanning target along the first emission angle, and then at least one ultrasonic wave may be emitted to the scanning target along the second emission angle. The beam, secondly, transmits at least one ultrasonic beam to the scanning target along the third emission angle to complete one scanning cycle, and finally repeats the scanning cycle process until the number of scanning times at all emission angles is completed.
  • Ultrasonic beam at different emission angles in the same scanning period The number of shots may be the same or different.
  • the transmitted ultrasonic beam is along two emission angles, then according to A1 B1 A2 B2 A3 B3 A4 B4 ... Ai Bi, and so on.
  • Ai is the ith transmission in the first transmission angle
  • Bi is the ith transmission in the second transmission angle.
  • the ultrasonic beam is emitted along three emission angles, then according to A1 B1 B1C1 A2 B2 B2C2 A3 B3 B3C3 ... Ai Bi BiCi, and so on.
  • Ai is the ith shot in the first launch angle
  • Bi is the ith shot in the second launch angle
  • Ci is the ith shot in the third launch angle.
  • the two ultrasonic beams may be alternately transmitted.
  • the step S100 includes: in step S101, the multi-focus ultrasonic beam is transmitted to the scanning target for Obtain image data of the reconstructed ultrasound image.
  • Step S102 transmitting a plurality of planar ultrasonic beams to the scanning target along one or more emission angles to acquire image data for calculating velocity vector information.
  • the process of emitting a focused ultrasound beam to the scanning target may be inserted during the transmission of the planar ultrasound beam to the scanning target.
  • the multi-focus ultrasonic beam transmitted to the scanning target is uniformly inserted into the process of performing the above-described step S102.
  • any of the alternate transmission modes that enable at least a portion of the plurality of planar ultrasonic beams transmitted to the scanning target to be alternately executed with at least a portion of the plurality of focused ultrasonic beams transmitted to the scanning target.
  • a focused ultrasonic beam can be used to obtain a better quality ultrasonic image; and a high real-time velocity vector information can be obtained by using a high planar beam frame rate, and in order to have better synchronization in data acquisition.
  • Sexuality using two types of ultrasonic-shaped alternating emission.
  • the receiving circuit 4 and the beam combining module 5 receive the echo of the above-mentioned transmitted ultrasonic beam, and perform beam combining to obtain an ultrasonic signal. For example, when an echo of the focused ultrasound beam is received, a focused ultrasound signal is obtained; when an echo of the planar ultrasound beam is received, a planar ultrasound signal is obtained, and so on. Which type of ultrasonic beam is emitted in step S100, and corresponding type of ultrasonic signal is generated corresponding to the echo of which type of ultrasonic beam is received.
  • the focused ultrasonic beam corresponds to the focused ultrasonic signal
  • the planar ultrasonic beam corresponds to the planar ultrasonic signal
  • the divergent ultrasonic beam corresponds to the divergent ultrasonic signal, and the like, and is not enumerated here.
  • the transmitting and receiving functions can be received when each of the array elements participating in the ultrasonic transmission or each of the array elements is time-divisionally implemented.
  • the echo of the ultrasonic beam emitted in step S100, or dividing the array element on the probe into the receiving portion and the transmitting portion, and then receiving each of the array elements or each of the array elements participating in the ultrasonic receiving, in the above step S100 The echo of the transmitted ultrasonic beam, and so on.
  • the ultrasonic beam When the ultrasonic beam is emitted along one emission angle in step S100, the echo of the ultrasonic beam from the emission angle is received, correspondingly obtaining a set of ultrasonic signals.
  • the ultrasonic beam is transmitted along the plurality of emission angles in step S100, a plurality of sets of ultrasonic signals corresponding to the plurality of emission angles are obtained corresponding to the echoes of the ultrasonic beams that receive the plurality of emission angles. Can receive different transmissions based on different emission angles Multiple sets of ultrasonic signals at an angle.
  • the set of ultrasonic signals includes a plurality of ultrasonic signals
  • the plurality of ultrasonic signals may correspond to receiving a plurality of echo signals of the plurality of ultrasonic beams emitted at each of the emission angles, wherein the transmission of the one ultrasonic beam corresponds to obtaining the ultrasonic signals.
  • step S100 multiple plane ultrasonic beams are respectively transmitted to the scanning target along a plurality of different emission angles, and then echoes of the plane ultrasonic beams corresponding to the plurality of emission angles are respectively received, and multiple groups of planes belonging to different emission angles are obtained.
  • each set of planar ultrasonic signals includes at least two planar ultrasonic signals, each of which is derived from an echo obtained by performing a step of transmitting an ultrasonic beam to the scanning target at a single emission angle.
  • an echo of the focused ultrasonic beam is received to obtain a plurality of focused ultrasonic signals.
  • the transmitting circuit 2 excites the probe 1 to emit an ultrasonic beam toward the scanning target along one or more emission angles in step S100
  • the ultrasonic beam from the scanning target can be received by adjusting the aperture position of the receiving array element on the probe. Waves, obtaining ultrasonic signals along different receiving angles, as ultrasonic signals of different angles obtained in step S100, as shown in FIG. 4 and related description.
  • the process of transmitting an ultrasonic beam to a scanning target along multiple emission angles is described in the foregoing.
  • step S100 when receiving an echo from the ultrasound beam on the scanning target, the aperture position of the receiving array element in the probe is adjusted to the first position for receiving the emission angle. Acquiring the ultrasonic beam to obtain a first set of ultrasonic signals belonging to the first receiving angle, adjusting the aperture of the receiving array element to the second position, and receiving the echo of the ultrasonic beam of the transmitting angle to obtain the second receiving angle.
  • the second set of ultrasonic signals is the same, so that multiple sets of ultrasonic signals are obtained based on different receiving angles.
  • the transmitting circuit 2 excites the probe 1 to transmit an ultrasonic beam to the scanning target, and receives echoes of the ultrasonic beam respectively at a plurality of different receiving angles to obtain ultrasonic signals of different components at different receiving angles. And an echo signal corresponding to a set of ultrasonic beams is received from the scan target at a receiving angle for subsequent beam synthesis, processing of the ultrasonic image data, and calculation of the blood flow velocity vector.
  • the echoes of the plurality of sets of ultrasonic beams are respectively received from the scanning target along a plurality of receiving angles.
  • a planar ultrasonic beam is transmitted to the scanning target, and an echo of the ultrasonic beam is received multiple times along a receiving angle to obtain a set of planar ultrasonic signals, wherein the set of planar ultrasonic signals includes multiple planar ultrasonic signals, along different
  • the receiving angle receives the echoes of the plurality of sets of planar ultrasonic beams, thereby obtaining a plurality of sets of planar ultrasonic signals belonging to different receiving angles.
  • Ultrasonic signals obtained based on a launch angle or a receive angle can also be used after The blood flow velocity vector is calculated in a subsequent step and/or an ultrasound image is obtained. For example, if a plane ultrasonic beam is transmitted to the scanning target along a transmission angle in step S100, the echo of the ultrasonic beam is received multiple times along a receiving angle to obtain a set of planar ultrasonic signals, and the set of planar ultrasonic signals includes multiple planes. Ultrasonic signal.
  • this embodiment can also be replaced with the other ultrasonic shapes described above.
  • an ultrasonic signal along one or more angles may be obtained in step S100, where the angle may include an emission angle or a reception angle.
  • a set of ultrasonic signals is obtained corresponding to a single emission angle or a receiving angle, and a plurality of sets of ultrasonic signals are obtained corresponding to different emission angles or reception angles, and each set of ultrasonic signals includes at least one ultrasonic signal obtained along a transmission angle or a reception angle.
  • An ultrasound image of at least a portion of the scan target can be acquired based on any one of the sets of ultrasonic signals or a combination of more than two sets of ultrasonic signals. Further, based on any one or a combination of two or more sets of the plurality of sets of ultrasonic signals, the blood flow velocity vector of the target point in the region of interest can be acquired.
  • step S100 in order to facilitate the calculation and improve the image display effect, the ultrasonic signals from the plurality of angles in the scanning target are obtained by the probe, and the ultrasonic signals of the plurality of angles belong to different receiving angles or emission angles, according to the super
  • the different angles corresponding to the wave signals are stored as at least one set of data frames associated with the angle. That is, the set of ultrasonic signals obtained as described above is stored as a set of data frames related to the angle, and the data frame set includes at least one frame of image data.
  • step S100 The deformation based on FIG. 5 in FIG. 12 includes in step S100:
  • Step S191 transmitting a plane ultrasonic beam with different emission angles to the scanning target through the probe
  • Step S192 transmitting a focused ultrasonic beam to the scanning target through the probe;
  • Step S193 receiving an echo of the planar ultrasonic beam, obtaining a planar ultrasonic signal belonging to different emission angles, as a part of the ultrasonic signal or the ultrasonic signal that needs to be obtained in step S100, for calculating the blood flow velocity vector in step S200, Increase the calculation speed of the blood flow velocity vector.
  • Step S194 receiving an echo of the focused ultrasonic beam to obtain a focused ultrasonic signal, which is a part of the ultrasonic signal or the ultrasonic signal that needs to be obtained in step S100, to obtain at least a part of the scanning target according to the focused ultrasonic signal in step S601. Ultrasound image.
  • the ultrasonic image quality obtained in this embodiment is higher.
  • step S200 in Fig. 5 the image processing module 7 obtains the blood flow velocity direction within the scan target based on the ultrasonic signal obtained in step S100.
  • step S200 the direction of the blood flow velocity corresponding to all the target points in the entire imaging region of the scan target may be first calculated, and then the direction of the blood flow velocity at the plurality of target points to be obtained may be selected.
  • the target point of the blood flow velocity direction may be first determined, and then the ultrasound image is acquired to calculate a blood flow velocity direction of the target point corresponding to the plurality of target points respectively.
  • the target point in step S200 may be that the user is in the region of interest
  • the input pixel point or pixel area may also be a plurality of discrete pixel points or pixel areas automatically generated by the system in the region of interest for determining the blood flow velocity direction or the blood flow velocity vector at a certain image coordinate or a block of image coordinates. Association point.
  • the blood flow velocity vector at the plurality of target points in the scan target may be obtained according to the ultrasonic signal, and the blood flow velocity vector includes a blood flow velocity value and a blood flow velocity direction. Then choose to extract the direction of blood flow velocity at multiple target points. Of course, it is also possible to calculate the blood flow velocity direction without calculating the blood flow velocity value.
  • the target point mentioned in step S200 may be the actual position selected for the determination, or may be calculated based on the blood flow velocity vector calculated at the previous time.
  • the target point mentioned in step S200 may be the actual position selected for the determination, or may be calculated based on the blood flow velocity vector calculated at the previous time.
  • the offset of the same spot between the two frames of images is used to obtain a blood flow velocity vector for calculating the target point. Specifically, it is shown in FIG. 6.
  • an ultrasonic signal is acquired in the manner described above, and the ultrasonic signal may include at least one set of ultrasonic signals.
  • a planar ultrasonic signal can be used to acquire an ultrasound image of a blood flow velocity vector of a calculation target point.
  • the planar ultrasonic beam propagates substantially throughout the imaging region. Therefore, generally, the primary planar beam echo signal obtained corresponding to the planar ultrasonic beam transmitted once is processed to obtain one frame of planar beam echo image data.
  • the ultrasonic image data of the scanning target obtained by performing corresponding processing on the plane beam echo signal obtained by the plane ultrasonic beam is referred to as a "planar beam echo image”.
  • a tracking area is selected in the first frame ultrasound image, which may include a target point where it is desired to obtain its velocity vector or velocity direction.
  • the tracking area can select a neighborhood of the target point or a data block containing the target point, see the smallest box on the left in Figure 6.
  • searching for an area corresponding to the tracking area in the second frame ultrasound image for example, searching for an area having the greatest similarity with the aforementioned tracking area as a tracking result area (see, lower position in the largest square on the right side in FIG. 6)
  • the second small box, the first small box in the upper position of the largest box on the right side in Fig. 6 indicates the position of the tracking area corresponding to the ultrasound image in the second frame).
  • the measure of similarity can use the following formula to find the similarity matrix, and based on the similarity matrix, the search has the most A region of great similarity.
  • the similarity matrix in the two-dimensional image is calculated by the following formula (1) or (2).
  • X 1 is a first frame ultrasound image and X 2 is a second frame ultrasound image.
  • i and j are the horizontal and vertical coordinates of the two-dimensional image. Indicates the value of K and L when the result of the expression on the right side of it reaches a minimum.
  • K, L represents the new position in the image.
  • M, N is the size of the tracking area in the figure. with Is the average of the first frame and the second frame tracking area and the tracking result area.
  • the similarity matrix in the three-dimensional image is calculated by the following formula (3) or (4).
  • X 1 is a first frame ultrasound image and X 2 is a second frame ultrasound image.
  • i, j and k are the coordinates of the three-dimensional image. Indicates the value of A, B, and C when the result of the expression on the right side of it reaches a minimum.
  • A, B, and C represent the new horizontal and vertical coordinate positions in the image.
  • M, N, L are the sizes of the tracking areas in the figure. with Is the average of the first frame and the second frame tracking area and the tracking result area.
  • the velocity vector of the target point can be obtained according to the foregoing tracking area and the position of the foregoing tracking result area, and the time interval between the first frame image data and the second frame image data.
  • the velocity value may pass the distance between the tracking area and the tracking result area (ie, the movement displacement of the target point within a preset time interval), divided by the first frame plane beam echo image data and the second frame plane beam echo
  • the time interval between the image data is obtained, and the speed direction may be the direction of the line from the tracking area to the tracking result area, that is, the moving direction of the target point within the preset time interval.
  • the blood flow velocity direction in step S200 can be obtained, and the blood flow velocity value can also be obtained, and the blood flow velocity vector can be obtained by integrating the blood flow velocity direction and the blood flow velocity value.
  • the obtained at least two frames of the ultrasonic image may be subjected to wall filtering processing, that is, wall filtering is performed on the points in each position on the image in the time direction.
  • wall filtering is performed on the points in each position on the image in the time direction.
  • the tissue signal on the image changes less with time, while the blood flow signal changes greatly due to the flow of blood. Therefore, Qualcomm can be used.
  • the filter acts as a wall filter for the blood flow signal. After wall filtering, the higher frequency blood flow signal is retained, and the smaller frequency tissue signal is filtered out. After the wall-filtered signal, the signal-to-noise ratio of the blood flow signal can be greatly enhanced.
  • the blood flow velocity vector of the target point is obtained based on the time gradient and the spatial gradient at the target point, as shown below.
  • an ultrasonic signal is acquired in the manner described above, and the ultrasonic signal may include at least one set of ultrasonic signals.
  • the ultrasonic signal may be an ultrasonic signal that is divided into one or more angles. This angle may be an emission angle or a reception angle.
  • the following embodiment illustrates the emission angle as an example.
  • the ultrasonic signal obtaining at least two frames of ultrasound images
  • obtaining a first gradient in the time direction at the target point according to the ultrasound image obtaining a second gradient along the emission angle at the target point according to the ultrasound image, obtaining a direction perpendicular to the emission angle at the target point according to the ultrasound image a third gradient, calculating a fifth velocity component of the target point at the emission angle and a sixth velocity component in a direction perpendicular to the emission angle according to the first gradient, the second gradient, and the third gradient;
  • the blood flow velocity vector of the target point is obtained according to the fifth speed component and the sixth velocity component, including the blood flow velocity value and the combined angle obtained after the synthesis, and the composite angle points to the blood flow velocity direction.
  • the emission angle employed in the above embodiment is taken as an embodiment. If at least two frames of the ultrasound image are obtained by using the reception angle as mentioned above, the calculation may be performed in the above manner, but the "emission" in each step. Angle should be replaced with the acceptance angle.
  • the above process uses a planar ultrasonic signal to perform calculations to improve the speed and accuracy of the velocity vector. Based on the above method, the blood flow velocity direction in step S200 can be obtained, and the blood flow velocity value can also be obtained, and the blood flow velocity vector can be obtained by integrating the blood flow velocity direction and the blood flow velocity value.
  • the blood flow velocity component along a plurality of different angles is obtained at the target point; the blood flow velocity component associated with the plurality of different angles is synthesized, and the target is obtained at the target point Blood flow velocity vector.
  • a Doppler imaging technique can be utilized to calculate a blood flow velocity component at an angle at a target point.
  • an ultrasonic signal is acquired in the manner described above, and the ultrasonic signal may be an ultrasonic signal belonging to a plurality of angles. This angle can be the launch angle or the receive angle.
  • the following embodiment will be described by taking an ultrasonic beam emitted to a scanning target along a plurality of emission angles and receiving an echo signal of the ultrasonic beam as an ultrasonic signal in step S100.
  • multiple ultrasonic beams are continuously transmitted at the same emission angle for the scanning target; the echoes of the multiple ultrasonic beams received are received, and multiple ultrasonic signals are obtained, and each value of each ultrasonic signal corresponds to The value at a target position when scanning at an emission angle.
  • step S200 the calculation is performed as follows:
  • V z is the calculated velocity value along the emission angle
  • c is the speed of sound
  • f 0 is the center frequency of the probe
  • T prf is the time interval between two shots
  • N is the number of shots
  • x(i) is The real part of the i-th shot
  • y(i) is the imaginary part of the ith shot.
  • the above formulas (5) and (6) are formulas for calculating the velocity values at a fixed position.
  • the velocity value at each target point can be found by the N complex values.
  • the Doppler velocity value V z can be used to characterize the blood flow velocity value along the corresponding emission angle at the target point, and the emission angle characterizes the blood flow velocity direction at the target point, and the combination obtains the edge.
  • the blood flow velocity component of the corresponding emission angle can be expressed in a vector manner.
  • the emission angle adopted in the above embodiment is taken as an embodiment. If a plurality of ultrasonic signals are obtained along a receiving angle as mentioned in the foregoing, the calculation may be performed in the above manner, but the transmission angle in the foregoing is replaced by receiving. Angle, the direction of blood flow velocity is the receiving angle, so that the blood flow velocity component along the corresponding receiving angle can be obtained.
  • the blood flow velocity values in different angular directions can be respectively obtained, which can be characterized by the Doppler frequency.
  • Doppler processing is performed on the ultrasonic signal by the Doppler principle, and the moving speed of the scanning target or the moving portion therein can be obtained.
  • the moving speed of the scanning target or the moving portion therein can be obtained from the ultrasonic signal by the autocorrelation estimation method or the cross-correlation estimation method.
  • the method of performing Doppler processing on the ultrasonic signal to obtain the velocity of motion of the scanning target or the moving portion therein may use any of the fields currently in use in the art or may be used in the future to calculate the scanning target or the ultrasonic signal therein. The method of the speed of the movement of the moving part will not be described in detail here.
  • the blood flow velocity components at different target angles or receiving angles at the target point are obtained, and then the velocity components are synthesized at the target point, and the composite velocity at the target point, that is, the blood flow velocity vector of the target point is obtained.
  • the composite velocity at the target point that is, the blood flow velocity vector of the target point is obtained.
  • At least two sets of ultrasonic signals are acquired in the manner described above, and the at least two sets of ultrasonic signals may be ultrasonic signals belonging to a plurality of different angles, and different angles include different emission angles or different reception angles (step S110).
  • the ultrasonic signals corresponding to different angles are stored as at least two sets of data frames related to the angle.
  • the ultrasonic signals corresponding to different angles are stored as at least two sets of data frames related to the angle.
  • the blood flow velocity component corresponding to each set of data frames is respectively calculated, and at least two blood related to the angle are obtained.
  • the flow velocity component (step S220). At least two blood flow velocity components are obtained at each target point.
  • step S200 the blood flow velocity vector desired to be obtained in step S200 is obtained, which includes the blood flow velocity value and the combined angle obtained after the synthesis, and the combined angle points to the blood flow velocity direction (step S230).
  • blood flow velocity vectors corresponding to the respective target points can be obtained.
  • the emission angle employed in the above embodiment is taken as an embodiment. If multiple sets of ultrasonic echo signals are obtained along a plurality of reception angles as mentioned in the foregoing, the calculation may be performed in the above manner, but in each step. The launch angle should be replaced by the "receiving angle".
  • the present invention is not limited to the above method for a blood flow velocity component corresponding to one emission angle or reception angle, and other methods known in the art or possible in the future may be employed.
  • the velocity values in the blood flow velocity vector may include approximate or true velocity, acceleration, velocity variance evaluation values, etc. at the target point, and statistics representing the velocity state. One of them.
  • step S120 or juxtaposed with step 120 further includes the following steps:
  • the image processing module obtains an ultrasound image of at least a portion of the scan target according to the ultrasonic signal, and the ultrasound image herein may be a three-dimensional ultrasound stereo image, or may be a two-dimensional ultrasound image, such as a B-picture, which is obtained by using the scan body for display.
  • the ultrasound image may be imaged using planar ultrasound beams or focused ultrasound beam imaging.
  • the ultrasound image can be imaged using a focused ultrasound beam.
  • step S100 The medium-transmitted multiple-shot focused ultrasound beam is used to scan to obtain a frame of ultrasound images.
  • the plurality of focused ultrasound beams are transmitted to the scanning target in the above step S100, and the echoes of the focused ultrasound beams are received in step S200 to acquire a set of focused beam echo signals according to the focus.
  • the beam echo signal obtains an ultrasound image of at least a portion of the scan target.
  • High quality ultrasound images can be obtained with focused ultrasound.
  • the data of acquiring the ultrasound image may be obtained based on any one of the ultrasonic signals or any set of data frames in the foregoing step S100.
  • the ultrasound image is displayed through the display and the sampling frame can also be displayed on the ultrasound image.
  • the sampling frame of this embodiment may be one or more, and multiple sampling frames may overlap. When sampling multiple sampling frames, you can simultaneously compare the quantitative results of multiple dispersions.
  • the sampling frame can be adjusted by the user to determine a redefined sampling frame based on the adjustment signal of the sampling frame by the user to obtain the size and shape of the sampling frame.
  • step S300 the image processing module 7 extracts a plurality of blood flow velocity directions.
  • the extracted multiple blood flow velocity directions include at least one of the following contents:
  • the same time includes the same time, or the same time period; the time includes at least one real time point; and the time period includes at least one time.
  • the time in this embodiment can also be determined by the acquisition frame rate of the image.
  • the corresponding blood flow velocity directions at a plurality of positions at the same time it is possible to evaluate the eddy currents at the same time at the same time, and to sample the blood flow velocity directions corresponding to the same position at different times, the same can be evaluated.
  • the extracted plurality of blood flow velocity directions include a plurality of blood flow velocity directions corresponding to any one of the same cardiac cycles, so that the same position in one cardiac cycle can be evaluated during different periods or during the same period.
  • the direction of blood flow changes.
  • the extracted plurality of blood flow velocity directions include: a plurality of blood flow velocity directions corresponding to the same time in different cardiac cycles, for quantitatively evaluating changes in the direction of blood flow velocity during the systolic or diastolic period.
  • the phase in this context includes any time or period of time in the cardiac cycle, including systolic and/or diastolic phases.
  • the location of this embodiment may be a point or region of interest within the scan target, typically represented as a point or region of interest in at least a portion of the ultrasound image of the scan target displayed on the display that may be marked or may be displayed.
  • the location described above can include at least one target point.
  • the direction of the blood flow velocity at the position may take the combined angular direction of the blood flow velocity vector corresponding to the plurality of target points at the position, or corresponding to the plurality of target points at the position One of the blood flow velocity directions, or the blood flow velocity corresponding to the plurality of target points at the position The most number of directions in the direction, and so on.
  • the blood flow velocity vector at the position may take the composite velocity of the blood flow velocity vector corresponding to the plurality of target points at the position, or take the blood flow velocity vector corresponding to the plurality of target points at the position. Or one of taking the highest number of directions in the blood flow velocity vector corresponding to the plurality of target points at the position, or taking the mean of the blood flow velocity vectors corresponding to the plurality of target points at the position, and the like.
  • the plurality of locations to be acquired in the direction of the blood flow velocity are determined according to the set region of interest, and the set region of interest may include the region of interest selected by the user, and automatically segmented based on the system image segmentation technique.
  • the obtained vascular area, the system selects the range of the region of interest and the entire imaging region of the scanning target, and the like, or a combination of one or more ranges thereof.
  • the plurality of locations may be a plurality of discrete or continuous locations of the region of interest, which may be automatically assigned by the system or selected by the user. It can be seen that, in step S300, multiple blood flow velocity directions in the region of interest are extracted to quantify the estimated region of interest.
  • the region of interest is determined by a sampling frame, which may be an area that the system automatically forms on the ultrasound image, or the entire imaging area, or may also be obtained by the user entering a selection instruction on the ultrasound image. Area, and so on.
  • the region of interest includes at least one target point, or a field (data block) containing at least one target point. For example, 31 in Fig. 13, 41 in Fig. 14, 51 or A32 in Fig. 15, 61 or 62 in Fig. 16, and 71 or 72 in Fig. 17.
  • the plurality of extracted blood flow velocity directions may be a combination of the first type and the second one.
  • the extracted plurality of blood flow velocity directions include a plurality of positions corresponding to blood flow velocity directions at a plurality of times. .
  • the above embodiment can comprehensively measure the flow of blood in the vessel from the spatial and temporal dimensions.
  • the image processing module performs step S600 to obtain an ultrasound image of at least a portion of the scan target according to the ultrasonic signal, and the display performs step S700 to display the ultrasound image, and performs step S800.
  • the operation control module acquires the sampling frame on the ultrasound image, and after step S200, the step S310 is performed, and the image processing module extracts a plurality of blood flow velocity directions associated with the sampling frame, and the plurality of blood flow velocity directions may be discrete multiples in the sampling frame.
  • the extracted plurality of blood flow velocity directions includes at least a portion of the position-dependent blood flow velocity directions within the sampling frame at the same time, and the sampling frame is time-dependent blood flow. At least a portion of the speed direction.
  • the plurality of extracted blood flow velocity directions may be a plurality of blood flow velocity directions corresponding to all positions in the sampling frame, or may be a plurality of blood flow velocity directions corresponding to the partial positions. In this way, it is possible to extract the positions in the sampling frame that are expected to be discretized, and the desired positions may be relatively close or distant, so that the macroscopically more flexible points can be divided. The dispersion of the direction of blood flow velocity over a relatively large range is analyzed.
  • the method further includes: recording a correspondence relationship between the blood flow velocity vector at each position and time; and then extracting more from the corresponding relationship when extracting the plurality of blood flow velocity directions The direction of blood flow velocity corresponding to each moment. Thereby, the time information about the direction of the blood flow velocity can be more clearly obtained, which is convenient for display in the subsequent process.
  • the direction of blood flow velocity is represented by an angle value or a direction vector.
  • the direction of blood flow velocity is represented by an angle value of 0-360 degrees, or by coordinates in spherical coordinates or Cartesian coordinates.
  • step S400 the image processing module quantizes the dispersion of the extracted plurality of blood flow velocity directions.
  • the degree of dispersion represents the degree of difference between the extracted multiple blood flow velocity directions. Quantization refers to the use of values to represent dispersion.
  • the dispersion of the plurality of blood flow velocity directions extracted by quantization is one of the following ways:
  • the variance of the plurality of blood flow velocity directions is calculated, the angular difference extreme values of the plurality of blood flow velocity directions are calculated, and the standard deviation of the plurality of blood flow velocity directions is calculated.
  • the specific calculation method is as follows.
  • the velocity direction of each point in space needs to be represented by two angle values, namely ⁇ and The speed is represented by r. It can also be represented by a vector in a Cartesian coordinate system by means of coordinate transformation, ie (x, y, z), which contains the magnitude and direction information of the velocity.
  • the dispersion of the velocity direction (angle) at a plurality of points can be calculated using the following vector variance or standard deviation.
  • the two-dimensional vector of the blood flow velocity vector is expressed as (x 1 , y 1 ), (x 2 , y 2 ), ... (x N , y N ), and the unit vector is obtained by processing. Etc., they only have blood flow velocity direction information, and (a, b) represents a two-dimensional unit vector representing the direction of blood flow velocity.
  • the two-dimensional vector variance Var2 used to quantify the dispersion of multiple blood flow velocity directions can be expressed as:
  • the two-dimensional vector standard deviation SD2 used to quantify the dispersion of multiple blood flow velocity directions is:
  • the three-dimensional vector of the blood flow velocity vector (x 1 , y 1 , z 1 ), (x 2 , y 2 , z 2 ), ... (x N , y N , z N ) is the same as above, Processing to get the unit vector Etc., they only have directional information, and (a, b, c) represent a three-dimensional unit vector representing the direction of blood flow velocity.
  • the three-dimensional vector variance Var3 used to quantify the dispersion of multiple blood flow velocity directions can be expressed as
  • the angular difference extreme values of the plurality of blood flow velocity directions are obtained in the following manner.
  • d i represents the magnitude of the angle of the i-th blood flow velocity direction
  • d j represents the magnitude of the angle value of the j-th blood flow velocity direction.
  • d i or d j They are an angular value of 0-360 degrees.
  • the angle difference extreme value calculated according to the above formulas (16) and (17) is a value varying between 0 and 180 degrees. The larger the value, the larger the extreme value of the angular difference.
  • the angle difference maximum can also be measured by distance, such as the following formulas (20) and (21)
  • the angular difference between any two angles is first calculated; then, the maximum or minimum value of the angular differences is searched for.
  • the angular difference in this embodiment can be expressed as in the formulas (18) and (19), or the distance measurement in the formulas (20) and (21).
  • the blood flow velocity direction is quantized to a value between 0 and 360 degrees (or -180 to 180 degrees), and the variance Var can be calculated using the following formula (22).
  • ⁇ in the formula is the angle value after the quantification of the blood flow velocity direction, and N represents the number of target points in the sampling frame or the number of positions of different sampling frames.
  • Var is a 0-1 number, you can also multiply it by 100, then the variance is a 0-100 number, and so on can be modified accordingly.
  • the larger the Var the greater the change in the velocity direction in the sampling frame (the greater the dispersion, the greater the degree of turbulence or eddy currents).
  • Laminar flow for example, when Var is 0, the blood flow velocity directions of all points in the sampling frame are completely the same, which is a typical laminar flow state.
  • the method of quantifying the dispersion of a plurality of blood flow velocity directions is not limited in the present invention, and other user-selected methods may also be adopted.
  • a variety of quantitative methods can be provided in the system for the user to select, and a comparative view of the dispersion results obtained by different quantization methods is performed to achieve comprehensive evaluation.
  • step S500 the quantized result of the dispersion is displayed by the display.
  • the display may also include displaying the ultrasound image with the display and marking the sampling frame in the ultrasound image for comparison to view the quantized result of the dispersion at the associated region.
  • the quantized result of the dispersion is superimposed on the ultrasound image.
  • the method of superimposing the quantized result of the dispersion on the ultrasound image includes:
  • an image processing module is utilized to generate a particle block associated with a particular region, the color coding of the particle block being associated with discrete quantized results of blood flow velocity directions within a particular region. Then, a color-coded particle block is displayed at a specific area of the ultrasound image to obtain a dispersion image. For the particle block, see the color spots in Figures 13 and 14 and the color block partitions of A1, A2, A3, and A4.
  • the specific area in this embodiment means that the blood flow velocity direction quantization result in a specific area on the ultrasonic image corresponds to one particle point block, and the specific area may be set according to a system setting or a user selected image area range, for example, 3 *8,4*4,5*5.
  • the ultrasound image is segmented according to the range of the specific region, and the discrete quantization results of the corresponding blood flow velocity directions in the plurality of specific regions in the ultrasound image are sequentially calculated, where the plurality of specific regions may be obtained by non-overlapping segmentation of the ultrasound image.
  • the area size of each specific area may be the same or different.
  • the ultrasound image is 80*80, and may be equally divided into 100 8*8 specific areas, and each specific area does not overlap.
  • the dispersion of the blood flow velocity direction of each specific region is sequentially calculated by the direction of the blood flow velocity of each target point in each specific region, and the calculation result is the discrete quantization result of the center point in the specific region.
  • the corresponding specific area is color-coded according to the size of the discrete quantization result and displayed on the ultrasound image.
  • a plurality of specific regions herein may also be the result of continuous overlapping segmentation of the ultrasound image.
  • a small sampling frame is passed through the entire sampling frame (see the large box in FIG. 8). (See the small box in Fig. 8) After shifting one pixel point or a plurality of pixel points one by one and then moving to the next position, a new image area range is obtained, and the ultrasonic images are successively overlapped and sequentially.
  • the small sampling frame continuously offsets the plurality of image region ranges obtained, corresponding to a plurality of specific regions.
  • the calculated value is the discrete quantization result at the corresponding center point position, and the corresponding small sampling frame is sequentially color coded according to the size of the discrete quantization result, and the dispersion image is formed on the ultrasound image.
  • the two sampling frames can have overlapping portions, and the more overlapping, the higher the spatial resolution of the calculated dispersion image.
  • FIG. 13 shows a dispersion image of a color effect in which a portion of a vessel within the sampling frame 31 superimposes a discrete quantization result of a corresponding specific region to obtain a dispersion image such as that having a plurality of color patches in FIG. Also, a color code Bar 32 is provided on the interface to prompt the user to identify the size of the dispersion.
  • FIG. 14 is a diagram showing the line effect of FIG. 13, the sampling frame is 41, and the ultrasonic image is 42, wherein A1, A2, A3, and A4 respectively display different colors to show the difference in dispersion in the current region. In addition, if the rendering is performed in the manner shown in FIG.
  • the patch partitions of A1, A2, A3, and A4 will be further subdivided, so that the spatial resolution of the discrete image is improved. It can be seen that a new imaging effect diagram is provided in the embodiment, which can provide a more intuitive observation angle to the doctor, and exhibits blood flow movement in the vessel from the superimposed display effect with the ultrasound image.
  • Fig. 13 green indicates the smallest dispersion, and red indicates the largest dispersion.
  • the quantized result of the dispersion may be displayed directly at any position on the display interface based on the manner of text display, for example, "SD 0.01/100" shown in FIGS. 16 and 17.
  • an icon model constructed based on the quantified results is constructed, and an icon model is displayed to show discrete quantized results of associated sample frames.
  • the icon model can be displayed. For example, as in Figure 642, a color-coded rectangular column is used, and the color coding is related to the quantization result size of the dispersion. Also for example, as in 641 of Figure 16, a circular icon with an arrow pointing is used, where the arrow pointing is related to the discrete quantized result.
  • an icon model based on the positional relationship. For example, as shown in FIG. 18, the sampling frame 81 in the display area 80 in the ultrasound image, or the identifier 83 in the region of interest 91 by displaying the blood flow velocity vector, is constructed.
  • An icon model 84 corresponding to a partial image area of the vessel or vessel, partitioning the icon modulo 84 (B1, B2, B3, B4), and calculating the dispersion of each partition corresponding to the corresponding area on the ultrasound image The result is quantized and then the discrete quantized result is superimposed on the corresponding partition in text or color coding for presentation on the display.
  • the icon model can also be constructed based on time variation, for example, constructing the relationship between the dispersion and the time, recording the variation of the quantized result of the dispersion with time, and displaying the variation of the quantitative result of the dispersion with time, in the relationship coordinates.
  • the quantized results of the dispersion are plotted one by one to produce a dispersion change map associated with the sampling frame.
  • the dispersion change graph belongs to one of the icon models. Referring to FIG.
  • the ultrasonic image 50 includes a large sampling frame 51 and a small sampling frame A32; corresponding to the dispersion quantization result chart 53 of the small sampling frame A32, the dispersion is established (for example, Variance SD) and The coordinate relationship of time, and then extract the discrete quantization results at each time t31, t32, t33, t34, t35, draw it in the coordinate relationship, and obtain the dispersion change graph shown in the chart 53 for displaying the small sampling frame. Quantification of the dispersion of A32 at different times.
  • the icon 53 is also marked with a black inverted triangle, indicating the position of the small sampling frame A32, and can also be used to indicate the position corresponding to the current time.
  • FIG. 19 also provides another display effect diagram in which the ultrasound image 90 includes a large sampling frame 91 and a small sampling frame 92, corresponding to the dispersion quantization result chart 93 of the small sampling frame 92, establishing a time axis. Coordinate relationship, draw the discrete quantization results corresponding to each time t31, t32, t33, t34, t35, etc. on the time axis coordinates, draw rectangular columns corresponding to different time on the time axis, the color or height of each rectangular column Correlating with the quantized result of the dispersion corresponding to the current time, a dispersion change map is formed.
  • the quantized result of the dispersion is correlated with the corresponding position range in the ultrasonic image, and when the position range is changed, the quantized result is also updated.
  • the target point or the target point area of interest in the ultrasound image is marked by the sampling frame, then when the adjustment signal of the sampling frame is obtained by the user, after determining the redefined sampling frame according to the adjustment signal, in the step S300 extracts a plurality of blood flow velocity directions associated with the redefined sampling frame, where the meaning associated with the sampling frame includes selecting a plurality of blood flow velocity directions in the sampling frame, or further comprising determining the extracted by the number of sampling frames The number of blood flow velocity directions.
  • the quantization result changes as the sampling frame is updated.
  • the change also includes that the number of quantized results and/or the value of the quantized result changes as the sampling frame is updated.
  • the discrete quantified results of the direction of blood flow velocity may be displayed separately, or may be displayed together with the ultrasound image on the display interface of the display, where the ultrasound image may be a Doppler blood flow image, a blood flow ejection image, a B image, etc. One of them. That is, when the quantized result of the dispersion is displayed, the ultrasonic image is simultaneously displayed, and the blood flow velocity vector is superimposed on the ultrasonic image.
  • a display manner of superimposing a blood flow velocity vector on an ultrasound image is provided below.
  • the step S200 includes: calculating, according to the ultrasonic signal obtained in the above step S100, a blood flow velocity vector at a first display position in the ultrasound image of the target point at different times, to obtain a target point located Blood flow velocity vector information in ultrasound images at different times. Then, in the following process, the contrast display may be blood flow velocity vector information at the first display position in the ultrasound image at each time.
  • the ultrasonic image data P1, P2, ..., Pn corresponding to the times t1, t2, ..., tn can be respectively obtained, and then the target point is calculated.
  • the blood flow velocity vector at the first display position (the position of the black dot in the figure) in the ultrasound image at each time.
  • the target point is always located at the position (H1, W1) in the two-dimensional image in the ultrasound image at each time. Based on this, when the blood flow velocity vector information is compared and displayed in the subsequent step S800, that is, the difference is displayed at the target point (H1, W1) in the ultrasonic image P0 displayed on the display. Corresponding to the calculated blood flow velocity vector. If the target point is selected according to the user's self-selection part or all, or by the system default, the corresponding first display position can be obtained by the corresponding point, and the first display position in the ultrasonic image corresponding to the current time is calculated.
  • FIG. 9( a ) shows an effect diagram of the two-dimensional image P0 when it is displayed, and of course, it can also be applied to the three-dimensional image display, that is, the ultrasound image at each moment is taken as the scan body mentioned above to obtain a three-dimensional image database, and The first display position is taken as a spatial three-dimensional coordinate position in the three-dimensional image database, and will not be described here.
  • the step S200 includes: calculating, according to the ultrasonic signal obtained in the above step S100, the blood flow velocity vector sequentially obtained by continuously moving the target point to the corresponding position in the ultrasonic image, thereby acquiring the blood of the target point.
  • Flow velocity vector In the present embodiment, the blood flow velocity vector moved from one position to another position of the ultrasonic image in a time interval by repeatedly calculating the target point to obtain the target point in the ultrasonic image after continuously moving from the initial position. Corresponding blood flow velocity vectors at respective respective locations. That is to say, the calculation position for determining the blood flow velocity vector in the ultrasonic image of the present embodiment can be obtained by calculation. Then, in the following process, the contrast display may be the blood flow velocity vector at the position obtained by calculation in the ultrasonic image at each time.
  • the ultrasonic image data P11, P12, ... corresponding to the times t1, t2, ..., tn can be respectively obtained.
  • the initial position of the target point is determined according to the partial or total selection of the target point by the user or the density of the system default target point, etc., as shown in FIG. 9(b) (H1, W1).
  • the first point of the calculation is then to calculate the blood flow velocity vector A1 in the ultrasound image P11 at the initial position at time t1.
  • the calculation target point i.e., the black dot in the figure
  • the position (H2, W2) on the ultrasonic image P12 at time t2 and then the ultrasonic image P12 is obtained based on the ultrasonic signal.
  • a target point at the first time t1 is found at the second display position on the ultrasonic image at the second time, and then the blood flow at the second display position is obtained according to the ultrasonic signal obtained in the above step S100.
  • the velocity vector thereby obtaining the blood flow velocity vector information of the target point at the time instant t2 in the ultrasound image P12.
  • the displacement of the two adjacent moments is obtained to obtain the displacement amount, and the target point is determined according to the displacement amount at the second moment.
  • the movement displacement of the target point at a time interval is calculated, and the corresponding position of the target point in the ultrasound image is determined according to the displacement, and the time is moved according to the time interval from the initially selected target point.
  • the interval may be determined by the system transmission frequency, or may be determined by the display frame rate, or may be a time interval input by the user, by calculating the position reached after the target point is moved according to the time interval input by the user, and then obtaining the position at the position.
  • Blood flow velocity vector information is used for comparison display.
  • N initial target points can be marked in the figure according to the manner described above, and each initial target point has an arrow to indicate the magnitude and direction of the flow velocity at this point, as shown in Fig. 9(b).
  • the corresponding blood flow velocity vector is obtained when the marker target point is continuously moved to the corresponding position, forming a logo that flows in time.
  • the original arrow of each point will change position, so that it can be formed by the movement of the arrow.
  • this display mode is referred to herein as the second mode, the same below.
  • the effect diagram of the two-dimensional image P10 is shown in the example of FIG.
  • the first display position is taken as a spatial three-dimensional coordinate position in the three-dimensional image database, and is not described here.
  • the superimposed display of the blood flow velocity vector further includes displaying the blood flow velocity when The blood flow velocity vector obtained in the above step S200 is subjected to slow-motion processing for comparing the blood flow velocity vector after the slow-release processing.
  • the blood flow velocity vector is first subjected to slow processing to generate a slow blood flow velocity vector; then, a slow blood flow velocity vector is superimposed on the ultrasound image to form the blood flow projection map, thereby realizing the blood flow velocity vector. A comparison with discrete quantified results.
  • the color code and/or length of the particle projectile is related to the blood flow velocity value at a particular location in the vessel by generating a particle projectile as a marker to depict a change in blood flow velocity at the target point; Transmitting the particle projecting body into a display, displaying a change of the particle projecting body over time at a specific position of the ultrasonic image for dynamically displaying blood flow in the blood vessel by dynamic display of the particle projecting body Exercise to obtain a blood flow projection map.
  • the particle projecting body further includes a direction indicator whose direction is related to the direction of the blood flow velocity.
  • the actual flow direction of the target point within the scan target can be clearly depicted in the displayed blood flow projection map, and the blood flow velocity of the current position changes with time is displayed compared to the corresponding display position only in the image.
  • the size and direction of the way the more accurate, more realistic and visual representation of the actual blood flow direction within the scan target.
  • the process of flowing blood flow can be described by flowing points or arrows, or other markers that can depict the direction. Referring to Figs. 15 to 19, the particle projecting body is indicated by an arrow 83.
  • the particle projecting body may also include only the direction indicator without carrying blood flow velocity value information, and the direction of the direction marker is related to the direction of blood flow velocity at a specific position in the scanning target.
  • a particle projecting body including a direction mark is displayed at a specific position of the ultrasonic image to dynamically exhibit a moving direction of blood flow in the scanning target.
  • the particle projecting body in this embodiment may be similar to the manner in which the arrow is expressed.
  • the length and/or thickness of the arrow may be used to express the blood flow velocity value, and the direction of the arrow may be used to express the direction of blood flow velocity.
  • the specific position in this embodiment means that the blood flow velocity vector displayed at a specific position on the ultrasonic image corresponds to one particle projecting body, and the specific position may be a position for marking the blood flow velocity value, for example, FIG. 9 ( a) and the first display position or the second display position mentioned in Fig. 9(b).
  • the plurality of blood flow velocity directions for calculating the dispersion includes: a plurality of blood flow velocity directions corresponding to any phase in the same cardiac cycle (eg, systolic phase and/or diastolic phase), or the same phase in different cardiac cycles ( For example, systolic or diastolic phases correspond to multiple blood flow velocity directions.
  • a plurality of temporal blood flow velocity directions within the cardiac cycle reference may be made to the embodiment shown in FIG. 17.
  • the ultrasound image 70 includes a large sampling frame 71 and a small sampling frame 72, which are selected using a small sampling frame 72.
  • the position to be viewed may be based on the electrocardiogram or the Doppler spectrum map 74 provided on the display interface, and the direction of the blood flow velocity corresponding to different time frames in the systole and/or diastolic phase may be located by the selection of the cursor 73. Then calculate a systolic or diastolic phase, or a discrete direction of blood flow in a cardiac cycle.
  • the graphic representing the cardiac cycle is displayed by using a display, including an electrocardiogram, a Doppler spectrogram, a video browsing axis including a multi-frame image of a cardiac cycle, and the like, and the graphic can be visually determined at any time within the cardiac cycle.
  • the user in the echocardiography mode, can be provided with an operation guide for the user to select.
  • FIG. 5 is a schematic flow chart of a parameter display method of an embodiment. It should be understood that although the various steps in the flowchart of FIG. 5 are sequentially displayed as indicated by the arrows, these steps are not necessarily performed in the order indicated by the arrows. Except as explicitly stated herein, the execution of these steps is not strictly limited, and may be performed in other sequences. Moreover, at least some of the steps in FIG. 5 may include a plurality of sub-steps or stages, which are not necessarily performed at the same time, but may be executed at different times, and the order of execution thereof is not necessarily This may be performed in sequence, but may be performed in parallel or alternately with other steps or at least a portion of the sub-steps or stages of the other steps. 10 to 12 are extended embodiments of FIG. 5, and related steps can be referred to the related descriptions.
  • the technical solution of the present invention which is essential or contributes to the prior art, may be embodied in the form of a software product carried on a non-transitory computer readable storage carrier (eg The ROM, the disk, the optical disk, and the server cloud space include instructions for causing a terminal device (which may be a mobile phone, a computer, a server, or a network device, etc.) to perform the methods described in various embodiments of the present invention.
  • a terminal device which may be a mobile phone, a computer, a server, or a network device, etc.
  • the direction of the blood flow is first calculated, and then the direction is further used to further evaluate the degree of eddy current or turbulence, thereby serving as a quantitative analysis method for judging the degree of stenosis.

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Abstract

一种超声血流的参数显示方法，其包括：通过探头（1）获得来自于扫描目标内的超声波信号（S100）；根据所述超声波信号，获得所述扫描目标内多个血流速度方向（S200）；提取多个血流速度方向（S300）；量化提取的多个血流速度方向的离散度（S400）；显示所述离散度的量化结果（S500）。其提供了一种量化评估血流运动方向的方法，并为用户提供了更好的观察视角。

Description

超声血流的参数显示方法及其超声成像系统 技术领域
本发明涉及超声系统中血流信息成像显示技术，特别是涉及一种超声血流的参数显示方法及其超声成像系统。
背景技术
在医学超声成像设备中，超声波辐射到被检查的物体之内，彩色多普勒血流仪与脉冲波和连续波多普勒一样，也是利用红细胞与超声波之间的多普勒效应实现显像的。彩色多普勒血流仪包括二维超声显像系统、脉冲多普勒(一维多普勒)血流分析系统、连续波多普勒血流测量系统和彩色多普勒(二维多普勒)血流显像系统。震荡器产生相差为π/2的两个正交信号，分别与多普勒血流信号相乘，其乘积经模/数(A/D)转换器转变成数字信号，经梳形滤波器滤波，去掉血管壁或瓣膜等产生的低频分量后，送入自相关器作自相关检测。由于每次取样都包含了许多个红细胞所产生的多普勒血流信息，因此经自相关检测后得到的是多个血流速度矢量的混合信号。把自相关检测结果送入速度计算器和方差计算器求得平均速度，连同经FFT处理后的血流频谱信息及二维图像信息一起存放在数字扫描转换器(DSC)中。最后，根据血流的方向和速度大小，由彩色处理器对血流资料作为伪彩色编码，送彩色显示器显示，从而完成彩色多普勒血流显像。
将频谱多普勒用于心脏瓣膜狭窄和动脉硬化病变等的定量诊断。血流在一个心动周期中的不同时刻，其方向可能有所不同。例如颈总动脉的血流正常情况下是层流，但如果出现斑块，发生动脉狭窄后，血流会变得较为紊乱。在收缩期时狭窄附近可能会产生涡流。涡流程度也是判断狭窄率的一项重要指标，通常采用涡流面积作为涡流程度的判断指标，但传统的彩色多普勒无法测量血流的方向，只能凭借红蓝两色以及相关临床经验手动描绘出涡流的面积，因此容易产生误差。涡流程度需要定量计算出具体的数值，才能使诊断更加可靠。
发明内容
基于此，有必要针对现有技术中的不足，提供一种超声血流的参数显示方法及其超声成像系统，提供了一种新的量化评估血流运动方向的方法，并为用户提供了更好的观察视角。
本发明的一个实施例中提供了一种超声血流的参数显示方法，其包括：
通过探头获得来自于扫描目标内的超声波信号；
根据所述超声波信号，获得所述扫描目标内的血流速度方向；
提取多个血流速度方向；
量化提取的多个血流速度方向的离散度；
显示所述离散度的量化结果。
本发明的一个实施例中提供了一种超声成像系统，其包括：
探头，用于向扫描目标发射超声波束；
接收电路和波束合成模块，用于接收来所述超声波束的回波信号，进行波束合成后获得超声波信号；
图像处理模块，用于根据所述超声波信号，获得所述扫描目标内的血流速度方向，提取多个血流速度方向，量化提取的多个血流速度方向的离散度；
显示器，用于显示所述离散度的量化结果。
附图说明
图1为本发明一个实施例的超声成像系统的框图示意图；
图2为本发明一个实施例的垂直发射的平面超声波束的示意图；
图3为本发明一个实施例的偏转发射的平面超声波束的示意图；
图4为本发明一个实施例中多角度接收的示意图；
图5为本发明一个实施例的方法流程示意图；
图6为本发明一个实施例中血流速度矢量的一种计算方法示意图；
图7为本发明一个实施例中血流速度方向的空间坐标表示示意图；
图8为本发明其中一个实施例中离散度图像的计算示意图；
图9(a)为本发明的其中一个实施例中第一模式下血流速度矢量信息计算示意图；
图9(b)为本发明的其中一个实施例中第二模式下血流速度矢量信息计算示意图；
图10为本发明一个实施例的方法流程示意图；
图11为本发明一个实施例的方法流程示意图；
图12为本发明一个实施例的方法流程示意图；
图13为本发明其中一个实施例的离散度图像的显示效果示意图；
图14为图13的线条示意图；
图15、图16、图17、图18和图19分别为本发明中多个实施例中超声图像和离散度量化结构的对比显示的示意图。
具体实施方式
图1为本发明一个实施例的超声成像系统的结构框图示意图。如图1所示，该超声成像系统通常包括：探头1、发射电路2、发射/接收选择开关3、接收电路4、波束合成模块5、信号处理模块6、图像处理模块7和显示器8。本文中的“多个”指大于等于2个。
在超声成像过程中，发射电路2将经过延迟聚焦的具有一定幅度和极性的发射脉冲通过发射/接收选择开关3发送到探头1。探头1受发射脉冲的激励，向扫描目标(例如，人体或者动物体内的器官、组织、血管等等，图中未示出)发射超声波，经一定延时后接收从目标区域反射回来的带有扫描目标的信息的超声回信号，并将此超声回波信号重新转换为电信号。接收电路接收探头1转换生成的电信号，并将这些超声回波信号送入波束合成模块5。波束合成模块5对超声波信号进行聚焦延时、加权和通道求和等处理，获得超声波信号，然后将超声波信号送入信号处理模块6进行相关的信号处理，例如滤波等。经过信号处理模块6处理的超声波信号送入图像处理模块7。图像处理模块7根据用户所需成像模式的不同，对信号进行不同的处理，获得不同模式的图像数据，然后经对数压缩、动态范围调整、数字扫描变换等处理形成不同模式的超声图像，如B图像，C图像，D图像等二维图像，此外，该超声图像还可包括三维图像。图像处理模块7生成的超声图像送入显示器8进行显示。此外，图像处理模块7还可以依据超声波信号计算扫描目标内目标点的血流速度矢量，可以通过渲染处理将计算获得的血流速度矢量渲染添加在超声图像输出给显示器进行显示，和/或将计算的血流速度矢量送给显示器进行有关速度信息的显示。图像处理模块7和信号处理模块6分离设置在不同的处理器上或者集成在同一个处理器9上。
本实施例中提到的目标点可以是超声图像上一个像素点或者包含至少两个像素点的区域块。目标点的血流速度矢量，包括速度值和速度方向，用于表征扫描目标内血流运动状态的流速信息。下文将会详细解释血流速度矢量的相关计算方式。
探头1通常包括多个阵元的阵列。在每次发射超声波或接收超声波时，探头1的所有阵元或者所有阵元中的一部分参与超声波的发射。此时，这些参与超声波发射的阵元中的每个阵元或者每部分阵元分别受到发射脉冲的激励，并分别发射超声波，这些阵元分别发射的超声波在传播过程中发生叠加，形成被发射到扫描目标的合成超声波束，该合成超声波束的传播方向即为本文中所提到的超声波的发射角度。参与超声波发射的阵元可以同时被发射脉冲激励；或者，参与超声波发射的阵元被发射脉冲激励的时间之间可以有一定的延时。通过控制参与超声 波的发射的阵元被发射脉冲激励的时间之间的延时，可改变上述合成超声波束的传播方向(即发射角度)，下文将具体说明。
通过控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，也可以使参与超声波的发射的各个阵元发射的超声波在传播过程中不会聚焦，也不会完全发散，而是形成整体上大体上为平面的平面波。或者，通过控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，可以使各个阵元发射的超声波束在预定位置叠加，使得在该预定位置处超声波的强度最大，也就是使各个阵元发射的超声波“聚焦”到该预定位置处，该聚焦的预定位置称为“焦点”，这样，获得的合成的超声波束是聚焦到该焦点处的波束，本文中称之为“聚焦超声波束”。发射聚焦超声波束的过程中，参与超声波的发射的阵元可以以预定的发射时延(即参与超声波的发射的阵元被发射脉冲激励的时间之间存在预定的时延)的方式工作，各阵元发射的超声波在焦点处聚焦，形成聚焦超声波束。又或者，通过控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，使参与超声波的发射的各个阵元发射的超声波在传播过程中发生发散，形成整体上大体上为发散波。本文中，称这种发散形式的超声波为“发散超声波束”。
线性排列的多个阵元同时给予电脉冲信号激励，各个阵元同时发射超声波，合成的超声波束的传播方向与阵元排列平面的法线方向一致。如图2所示的垂直发射的平面波，此时参与超声波的发射的各个阵元之间没有时延(即各阵元被发射脉冲激励的时间之间没有时延)，各个阵元被发射脉冲同时激励。生成的超声波束为平面波，即平面超声波束，并且该平面超声波束的传播方向与探头1的发射出超声波的表面大体垂直，即合成的超声波束的传播方向与阵元排列平面的法线方向之间的角度为零度。但是，如果施加到各个阵元间的激励脉冲有一个时间延时，各个阵元也依次按照此时间延时发射超声波束，则合成的超声波束的传播方向与阵元排列平面的法线方向就具有一定的角度，即为合成波束的发射角度，改变上述时间延时，也就可以调整合成波束的发射角度的大小和在合成波束的扫描平面内相对于阵元排列平面的法线方向的发射方向。例如，图3所示为偏转发射的平面波，此时参与超声波的发射的各个阵元之间有预定的时延(即各阵元被发射脉冲激励的时间之间有预定的时延)，各个阵元被发射脉冲按照预定的顺序激励。生成的超声波束为平面波，即平面超声波束，并且该平面超声波束的传播方向与探头1的阵元排列平面的法线方向成一定的角度(例如，图3中的角a)，该角度即为该平面超声波束的发射角度。通过改变时延时间，可以调整角a的大小。同理，无论是平面超声波束、聚焦超声波波束还是发散超声波束，均可以通过调整控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，来调整 合成波束的方向与阵元排列平面的法线方向之间所形成的合成波束的“发射角度”，这里的合成波束可以为上文提到的平面超声波束、聚焦超声波波束或发散超声波束等等。
此外，参照前文所述，对于二维超声换能器，可以理解为多个线性阵列的组合，因此，二维超声换能器也可以通过控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，来调整合成波束与阵元排列平面的法线方向之间所形成的合成波束的“发射角度”。
更进一步地，通过控制参与超声波的接收的阵元(本文简称接收阵元)的孔径位置，可以调整接收的超声波信号的接收角度。例如，如图4所示，超声探头包括阵元部1、阵元部2、阵元部3和阵元部4。阵元部1、阵元部2、阵元部3和阵元部4可以是一个阵元或者多个阵元。可以用阵元部1、阵元部2、阵元部3和阵元部4中的一个或多个的组合来作为接收阵元。图4中，向包含目标点位置A的扫描目标发射一个发射角度的超声波束时，利用阵元部1作为接收阵元，接收从扫描目标内某一个目标点位置A反射回来的超声波束的回波，根据阵元部1的孔径位置和目标点位置A的连线(图4中标记为实线)，可以确定当前时刻接收的超声波束的回波的接收角度a1。同时，根据利用阵元部2作为接收阵元，接收从扫描目标内某一个目标点位置A反射回来的超声波束的回波，根据阵元部2的孔径位置和目标点位置A的连线(图4中标记为虚线)，可以确定当前时刻接收的超声波束的回波的接收角度a2。从同一个目标位置A返回的超声波束的回波可以获得两个不同接收角度的超声波束的回波。因此，根据接收阵元的孔径位置和目标点位置之间的连线、与超声阵元排列平面的法线方向之间的夹角，来定义上述超声波束的回波的“接收角度”。通过改变探头上接收阵元的孔径位置，则可以改变超声波束的回波的“接收角度”，从而获得从扫描目标返回的不同接收角度的超声波信号。
基于上述解释，向扫描目标发射超声波束，期望从扫描目标获得多个角度的超声波信号时，既可以通过改变探头上接收阵元的孔径位置来改变超声波束的回波的接收角度，从而获得来自于扫描目标的不同接收角度对应的超声波信号；也可以通过控制参与超声波的发射的阵元被发射脉冲激励的时间之间的延时，改变超声波束的发射角度，基于不同发射角度的超声波束的回波，获得来自于扫描目标的不同发射角度对应的超声波信号。图像处理模块7可以根据不同角度的超声波信号，计算扫描目标内或者扫描目标中感兴趣区域内多个目标点的血流速度矢量。
此外，图1所示的超声成像系统中还包括操作控制模块10，用于接收来自操 作用户输入的调节信号，该调节信号包括对超声波束的发射角度、接收角度、超声波束类型等成像参数进行的调整，或者还可以包括对组织图像处理模块的图像、感兴趣区域或血流速度矢量的计算结果进行的调整。操作控制模块10可以为人机交互接口，例如键盘、滚轮、与带触摸功能的显示屏连接的触摸手势接收和计算模块、鼠标、有关手势控制信号的收发模块等等。图1中的显示器8包括一个或多个显示屏，本实施例中的显示屏可以为触摸显示屏、LED显示屏等等。
图像处理模块输出的图像数据或者量化结果还可以通过无线传输模块，传输到远端显示器上进行显示，本实施例的方案不限于台式超声设备，还可以包括纳入医疗互联网系统内的所有可用于展现超声图像的设备。
图5中提供了一种超声血流的参数显示方法，其提供了一种可对血管内血流出现涡流或者紊流程度进行评估的方式，可以作为更加直观的判断血管狭窄程度的定量分析方法。以下将详细说明。
图5中的步骤S100，利用接收电路4和波束合成模块5，通过探头1获得来自于扫描目标内的接收超声波信号。
在其中一些实施例中，利用发射电路2激励探头1向扫描目标发射超声波束，接收该超声波束的回波获得步骤S100中提及的超声波信号。本实施例中向扫描目标发射的超声波束可以包括：聚焦超声波束和非聚焦超声波束，其中非聚焦超声波束包括：虚源超声波束、非衍射超声波束、发散超声波束或平面超声波束等多种波束类型中的至少一种或者至少两种以上的组合(这里的“以上”包括本数，以下同)。当然，本发明的实施例中不限于以上几种类型的超声波束。可见步骤S100中的超声波信号可以是超声波束的回波信号。
在其中一些实施例中，在步骤S100中包括：步骤121：向扫描目标发射聚焦超声波束，接收聚焦超声波束的回波，获得聚焦超声波信号，用以重建超声图像、或计算血流速度矢量等。或者，在步骤S100中包括步骤122：向扫描目标发射平面超声波束，接收平面超声波束的回波，获得平面超声波信号，用以重建超声图像、或计算血流速度矢量等。又或者，在步骤S100中包括上述步骤121和步骤122：向扫描目标发射聚焦超声波束，用以获得聚焦超声波信号；向扫描目标发射平面超声波束，用以获得平面超声波信号。聚焦超声波信号可用作重建扫描目标的至少一部分超声图像，以求获取质量较好的超声图像作为背景图像，而在图5中的步骤S200中平面超声波信号还可以用作计算血流速度矢量的图像数据基础。
在步骤S100中若采用两种波束类型进行发射，则向扫描目标交替发射两种超声波束。例如，在向扫描目标发射平面超声波束的过程中插入向扫描目标发射 聚焦超声波束的过程，即，交替执行上述步骤121和步骤122。这样可以保证两种超声波束图像数据获取的同步性，提高多波束角度发射获得的血流速度矢量的精确度。
除了波束类型可以自由选择，在步骤S100中还可以接收多个角度的超声波信号用以计算血流速度矢量或超声图像。例如，在步骤S100中可以向扫描目标发射不同发射角度的超声波束，用以接收获得多个发射角度对应的超声波信号。或者，从扫描目标上接收不同接收角度对应的超声波信号。可见，多个角度的超声波信号可以对应于多个发射角度，或者多个接收角度。具体如下所示：
1、针对向扫描目标发射不同发射角度的超声波束，可以沿不同的发射角度接收多个角度的超声波信号。
在一些实施例中，在步骤S100中可以包括：沿多个发射角度向扫描目标发射超声波束，接收所述超声波束的回波获得多个发射角度对应的超声波信号，作为步骤S100中接收获得的多个角度的超声波信号。
在其中一个实施例中，在步骤S100中包括：沿多个发射角度向扫描目标发射超声波束，在该过程中，可以按照发射角度的不同交替执行向扫描目标发射超声波束的过程。例如，若沿两个发射角度向扫描目标发射超声波束，则先沿第一个发射角度向扫描目标发射超声波束，然后再沿第二个发射角度向扫描目标发射超声波束，完成一个扫描周期，最后依次重复上述扫描周期过程。或者，还可以先沿一个发射角度向扫描目标发射超声波束，再沿另一个发射角度向扫描目标发射超声波束，依次执行完所有发射角度后完成扫描过程。为获取不同的发射角度，可通过改变参与超声波发射的阵元中的每个阵元或者每部分阵元的时延来获得，具体可参照图2或图3的解释。
在其中一个实施例中，沿每个发射角度向扫描目标发射多次超声波束，用以获得多次超声波信号，供后续超声图像数据的处理。例如，沿多个发射角度分别向扫描目标发射多次非聚焦超声波束、或者沿多个发射角度分别向扫描目标发射多次聚焦超声波束。而每一次超声波束的发射对应获得一次超声波信号。
按照发射角度的不同交替执行向扫描目标发射多次超声波束的过程，能使获得的回波数据近似计算在同一时刻的目标点的血流速度矢量，提高速度矢量信息的计算精度。例如，若沿三个发射角度分别向扫描目标发射N次超声波束，可以先沿第一个发射角度向扫描目标发射至少一次超声波束，然后再沿第二个发射角度向扫描目标发射至少一次超声波束，其次再沿第三个发射角度向扫描目标发射至少一次超声波束，完成一个扫描周期，最后依次重复上述扫描周期过程直至完成所有发射角度上的扫描次数。同一个扫描周期内不同发射角度下的超声波束的 发射次数可以相同，也可以不相同。例如，如果是沿两个发射角度的发射超声波束，则按照A1 B1 A2 B2 A3 B3 A4 B4……Ai Bi，以此类推。其中，Ai是第一个发射角度中的第i次发射；Bi是第二个发射角度中的第i次发射。而如果是沿三个发射角度的发射超声波束，则按照A1 B1 B1C1 A2 B2 B2C2 A3 B3 B3C3……Ai Bi BiCi，以此类推。其中Ai是第一个发射角度中的第i次发射；Bi是第二个发射角度中的第i次发射；Ci是第三个发射角度中的第i次发射。
当上述步骤S100中选择向扫描目标发射两种波束类型的超声波束时，可以交替发射两种的超声波束，例如，上述步骤S100包括：步骤S101，向扫描目标发射多次聚焦超声波束，用以获取重建超声图像的图像数据。步骤S102，沿一个或多个发射角度向扫描目标发射多次平面超声波束，用以获取计算速度矢量信息的图像数据。然而，可以在向扫描目标发射平面超声波束的过程中插入向扫描目标发射聚焦超声波束的过程。比如，将向扫描目标发射的多次聚焦超声波束均匀插入到执行上述步骤S102的过程中。或者还可以，采用任何一种能实现上述向扫描目标发射多次平面超声波束的至少一部分与上述向扫描目标发射多次聚焦超声波束的至少一部分交替执行方案的任何一种交替发射方式。本实施例中可以利用聚焦超声波束获得质量较好的超声图像；而可以利用平面超声波束帧率高的特点获得高实时性的速度矢量信息，而且为了在数据获取上两者具有更好的同步性，采用两种类型的超声波形交替发射的方式。
接收电路4和波束合成模块5接收上述发射的超声波束的回波，进行波束合成获得超声波信号。例如，当接收聚焦超声波束的回波，则获得聚焦超声波信号；当接收平面超声波束的回波，则获得平面超声波信号，依次类推。在步骤S100中发射何种类型的超声波束，那么对应接收何种类型的超声波束的回波，生成对应类型的超声波信号。例如，聚焦超声波束对应聚焦超声波信号、平面超声波束对应平面超声波信号、发散超声波束对应发散超声波信号等等，在此不逐一列举。
接收电路4和波束合成模块5接收上述步骤S100发射的超声波束的回波时，可以利用参与超声波发射的阵元中的每个阵元或者每部分阵元分时实现发射和接收功能时接收上述步骤S100中发射的超声波束的回波，或者将探头上的阵元分为接收部分和发射部分、然后利用参与超声波接收的阵元中的每个阵元或者每部分阵元接收上述步骤S100中发射的超声波束的回波，等等。
当在步骤S100中沿一个发射角度上发射超声波束时，接收来自该发射角度的超声波束的回波，对应获得一组超声波信号。当在步骤S100中沿多个发射角度上发射超声波束时，对应接收多个发射角度的超声波束的回波，获得对应于多个发射角度的多组超声波信号。基于不同的发射角度，可以接收对应于不同发射 角度的多组超声波信号。此外，一组超声波信号包括多次超声波信号，多次超声波信号可以对应于接收沿每个发射角度上发射的多次超声波束的多次回波信号，其中一次超声波束的发射对应获得一次超声波信号。例如，在步骤S100中沿多个不同发射角度分别向扫描目标发射多次平面超声波束，则分别接收上述多个发射角度对应的平面超声波束的回波，获得分属于不同发射角度的多组平面超声波信号，其中每组平面超声波信号包括至少两次平面超声波信号，每次平面超声波信号源自沿一个发射角度上执行一次向扫描目标发射超声波束的步骤所获得的回波。又例如，对于步骤S100中向扫描目标发射多次聚焦超声波束，则接收上述聚焦超声波束的回波，获得多次聚焦超声波信号。
2、沿不同的接收角度从扫描目标接收多个角度的超声波信号。
当在步骤S100中发射电路2激励探头1沿一个或多个发射角度向扫描目标发射超声波束时，可以通过调节探头上的接收阵元的孔径位置，来接收来自于扫描目标的超声波束的回波，获得沿不同接收角度的超声波信号，作为步骤S100中接收获得的不同角度的超声波信号，具体可参见图4及相关说明所示。沿多个发射角度向扫描目标发射超声波束的过程参见前文相关说明。
例如，在其中一个实施例中，在步骤S100中，当接收来自扫描目标上超声波束的回波时，则将探头中接收阵元的孔径位置调整到第一位置，用于接收该发射角度的超声波束的回波，获得属于第一接收角度的第一组超声波信号，将接收阵元的孔径调整到第二位置，用于接收该发射角度的超声波束的回波，获得属于第二接收角度的第二组超声波信号，同理，从而基于不同的接收角度获得多组超声波信号。
参照前文沿多个发射角度的执行顺序和规则，在上述实施例中沿不同的接收角度从扫描目标接收多个角度的超声波信号的过程中，也可以按照接收角度的不同交替执行多组超声波信号的接收过程。在本发明的其中一个实施例中，发射电路2激励探头1向扫描目标发射超声波束，分多个不同的接收角度分别接收该超声波束的回波，获得多组分属不同接收角度的超声波信号，其中沿一个接收角度从扫描目标对应接收一组超声波束的回波信号，供后续波束合成、超声图像数据的处理和血流速度矢量的计算。沿多个接收角度分别从扫描目标接收多组超声波束的回波。例如，在步骤S100中向扫描目标发射平面超声波束，沿一个接收角度多次接收超声波束的回波，获得一组平面超声波信号，此一组平面超声波信号中包括多次平面超声波信号，沿不同接收角度接收多组平面超声波束的回波，从而获得分属不同接收角度的多组平面超声波信号。
3、基于一个发射角度或者一个接收角度获得的超声波信号，也可以用于后 续步骤中计算血流速度矢量和/或获得超声图像。例如，在步骤S100中沿一个发射角度向扫描目标发射平面超声波束，则沿一个接收角度多次接收超声波束的回波，获得一组平面超声波信号，此一组平面超声波信号中包括多次平面超声波信号。当然此实施例中还可以替换为上述其他超声波形。
基于前文中提到的调整发射角度或者调整接收角度，均可以在步骤S100中获得了沿一个角度或多个角度的超声波信号，此处的角度可以包括发射角度或者接收角度。与一个发射角度或者接收角度对应获得一组超声波信号，对应不同的发射角度或接收角度可以获得多组超声波信号，每一组超声波信号中包括至少一次沿发射角度或接收角度获得的超声波信号。依据其中任意一组超声波信号或两组以上的超声波信号的组合，可以获取扫描目标的至少一部分的超声图像。此外，基于多组超声波信号中的任意一组或两组以上的组合，可以获取感兴趣区域中目标点的血流速度矢量。
在步骤S100中，为了便于计算方便和提升图像显示效果，通过探头获得来自于扫描目标内的多个角度的超声波信号中，多个角度的超声波信号分属于不同的接收角度或发射角度，按照超波信号对应的不同角度，存储为与角度相关的至少一组数据帧集。也就是将上述获得的一组超声波信号存储为与角度相关的一组数据帧集，数据帧集中包括至少一帧图像数据。
图12中基于图5的变形，在步骤S100中包括：
步骤S191，通过探头向扫描目标发射不同发射角度的平面超声波束，
步骤S192，通过探头向扫描目标发射聚焦超声波束；
步骤S193，接收所述平面超声波束的回波，获得分属于不同发射角度的平面超声波信号，作为步骤S100需要获得的超声波信号或超声波信号的一部分，用于在步骤S200中计算血流速度矢量，提升血流速度矢量的计算速度。
步骤S194，接收所述聚焦超声波束的回波，获得聚焦超声波信号，作为步骤S100需要获得的超声波信号或超声波信号的一部分，用以在步骤S601中根据聚焦超声波信号，获得扫描目标的至少一部分的超声图像。本实施例获得超声图像质量更高。
图5中的步骤S200中，图像处理模块7根据步骤S100中获得的超声波信号，获得扫描目标内的血流速度方向。
步骤S200中可以先计算扫描目标的整个成像区域中所有目标点对应的血流速度方向，然后选择待获得多个目标点处的血流速度方向。或者，还可以先确定待获取血流速度方向的目标点，然后获取超声图像用以计算此多个目标点分别对应的目标点的血流速度方向。步骤S200中的目标点可以是用户在感兴趣区域中 输入的像素点或像素区域，也可以是感兴趣区域中系统自动生成的多个离散像素点或像素区域，用于确定计算某个或某块图像坐标处血流速度方向或血流速度矢量的关联点。
在获得血流速度方向之前，可以先根据超声波信号，获得扫描目标内多个目标点处的血流速度矢量，该血流速度矢量包括血流速度值和血流速度方向。然后再选择提取多个目标点处的血流速度方向。当然，也可以不用计算血流速度值，而在计算获得血流速度方向。
此外，基于后续在超声图像上叠加血流速度矢量的显示方式不同，步骤S200中提到的目标点可以是选择确定的实际位置，也可以是根据前一时刻计算的血流速度矢量演算出来的位置，具体参见下文有关第一显示模式和第二显示模式的相关说明。
无论是血流速度方向的获取，还是血流速度值的获取，均可以参照以下几种计算方式。
第一种，基于斑点追踪的方式利用两帧图像之间相同斑点的偏移，来获得计算获得目标点的血流速度矢量。具体结合图6所示。
首先，按照前文所述方式获取超声波信号，此超声波信号可以包括至少一组超声波信号。
其次，根据所述超声波信号，获得至少两帧超声图像，例如获得至少第一帧超声图像(参见，图6中左边最大的方框)和第二帧超声图像(参见，图6中右边最大的方框)。如前文所述，本实施例中可以采用平面超声波信号来获取计算目标点的血流速度矢量的超声图像。平面超声波束大体上在整个成像区域中传播，因此，通常，一次发射的平面超声波束所对应获得的一次平面波束回波信号通过处理即可获得一帧平面波束回波图像数据。本文中，将对平面超声波束对应获得的平面波束回波信号进行相应的处理而获得的扫描目标的超声图像数据称之为“平面波束回波图像”。
然后，在第一帧超声图像中选择跟踪区域，该跟踪区域可以包含希望获得其速度矢量或速度方向的目标点。例如，跟踪区域可以选择目标点的某个邻域或者包含目标点的某个数据块，参见图6中左边最小的方框。
其次，在第二帧超声图像中搜索与该跟踪区域对应的区域，例如，搜索与前述的跟踪区域具有最大相似性的区域作为跟踪结果区域(参见，图6中右边最大方框中靠下方位置的第二小方框，图6中右边最大方框中靠上方位置的第一小方框表示跟踪区域对应在第二帧超声图像中的位置)。这里，相似性的度量过程可以采用下述公式来寻找相似矩阵，基于相似矩阵来寻找与前述的跟踪区域具有最 大相似性的区域。
二维图像中相似矩阵采用下述公式(1)或(2)计算。
Figure PCTCN2016101208-appb-000001
Figure PCTCN2016101208-appb-000002
其中，X1为第一帧超声图像，X2为第二帧超声图像。i和j是二维图像的横纵坐标。
Figure PCTCN2016101208-appb-000003
表示当它右边的式子计算结果达到最小时，K和L的值。K，L则代表图像中新的位置。M，N为图中跟踪区域的大小。
Figure PCTCN2016101208-appb-000004
和
Figure PCTCN2016101208-appb-000005
是第一帧和第二帧跟踪区域和跟踪结果区域中的平均值。
三维图像中相似矩阵采用下述公式(3)或(4)计算。
Figure PCTCN2016101208-appb-000006
Figure PCTCN2016101208-appb-000007
其中，X1为第一帧超声图像，X2为第二帧超声图像。i，j和k是三维图像的坐标。
Figure PCTCN2016101208-appb-000008
表示当它右边的式子计算结果达到最小时，A，B，C的值。A，B，C则代表图像中新的横纵坐标位置。M，N，L为图中跟踪区域的大小。
Figure PCTCN2016101208-appb-000009
和
Figure PCTCN2016101208-appb-000010
是第一帧和第二帧跟踪区域和跟踪结果区域中的平均值。
最后，根据前述的跟踪区域和前述的跟踪结果区域的位置，以及第一帧图像数据与第二帧图像数据之间的时间间隔，即可获得所述目标点的速度矢量。例如，速度值可以通过跟踪区域和跟踪结果区域之间的距离(即目标点在预设时间间隔内的移动位移)、除以第一帧平面波束回波图像数据与第二帧平面波束回波图像数据之间的时间间隔获得，而速度方向可以为从跟踪区域到跟踪结果区域的连线的方向，即目标点在预设时间间隔内的移动方向。
基于上述方法，可以获得步骤S200中的血流速度方向，也可以获得血流速度值，综合血流速度方向和血流速度值可以获得血流速度矢量。
此外，在进行速度计算前，还可以对获得的至少两帧超声图像进行壁滤波处理，就是对于图像上每个位置上的点沿时间方向分别做壁滤波。图像上的组织信号随时间变化较小，而血流信号由于血流的流动则变化较大。因此可以采用高通 滤波器作为血流信号的壁滤波器。经过壁滤波之后，频率较大的血流信号保留下来，而频率较小的组织信号将被滤去。经过壁滤波后的信号，血流信号的信噪比可大大增强。
第二种，基于目标点处的时间梯度和空间梯度获得目标点的血流速度矢量，具体如下所示。
首先，按照前文所述方式获取超声波信号，此超声波信号可以包括至少一组超声波信号。此超声波信号可以是分属一个或多个角度的超声波信号。此角度可以为发射角度或接收角度，以下实施例以发射角度为例说明。
其次，根据超声波信号，获得至少两帧超声图像；
然后，根据超声图像获得在目标点处沿时间方向的第一梯度，根据超声图像获得在目标点处沿发射角度的第二梯度，根据超声图像获得在目标点处沿垂直于发射角度的方向的第三梯度，根据第一梯度、第二梯度和第三梯度计算目标点在发射角度上的第五速度分量和在垂直于发射角度的方向上的第六速度分量；
其次，根据第五速度分量和第六速度分量合成获得目标点的血流速度矢量，其中包括合成后获得的血流速度值和合成角度，合成角度指向血流速度方向。
在上述实施例中采用的发射角度作为实施例，如果至少两帧超声图像采用前文中提到的沿接收角度获得超声波信号，则也可以采用上述方式来进行计算，但是每个步骤中的“发射角度”应被替换为接收角度。在其中一个实施例中，上述过程采用平面超声波信号来进行计算可以提升速度矢量的计算速度和精确度。基于上述方法，可以获得步骤S200中的血流速度方向，也可以获得血流速度值，综合血流速度方向和血流速度值可以获得血流速度矢量。
第三种，基于分属不同角度的数据帧集，在目标点处关联获得沿多个不同角度的血流速度分量；合成与多个不同角度相关的血流速度分量，获得该目标点处的血流速度矢量。
在其中一个实施例中，可以利用多普勒成像技术来计算在目标点处沿一个角度的血流速度分量。
首先，按照前文所述方式获取超声波信号，此超声波信号可以是分属多个角度的超声波信号。此角度可以为发射角度或接收角度。以下实施例以沿多个发射角度向扫描目标发射超声波束，并接受所述超声波束的回波信号作为步骤S100中的超声波信号为例进行说明。在多普勒超声成像方法中，针对扫描目标在同一发射角度连续发射多次超声波束；接收发射的多次超声波束的回波，获得多次超声波信号，每一次超声波信号中每个值对应了在一个发射角度上进行扫描时一个目标位置上的值。
然后，在步骤S200中按照以下方式进行计算：
将对应于一个发射角度的一组超声波信号中的多次超声波信号分别沿发射角度所在的方向做Hilbert变换，得到采用复数表示每个目标点上值的多个图像数据；N次发射接收后，在每一个目标点上就有沿时间变化的N个复数值，然后，按照下述两个公式(5)和(6)计算目标点z在发射角度方向上的速度大小：
Figure PCTCN2016101208-appb-000011
Figure PCTCN2016101208-appb-000012
其中，Vz是计算出来的沿发射角度的速度值，c是声速，f0是探头的中心频率，Tprf是两次发射之间的时间间隔，N为发射的次数，x(i)是第i次发射上的实部，y(i)是第i次发射上的虚部，
Figure PCTCN2016101208-appb-000013
为取虚部算子，
Figure PCTCN2016101208-appb-000014
为取实部算子。以上公式(5)和(6)为一个固定位置上速度值的计算公式。
其次，以此类推，每个目标点上的速度值通过这N个复数值都可以求出。
如果采用上述方法来计算血流速度分量，那么可以取多普勒速度值Vz表征目标点上沿相应发射角度的血流速度值，发射角度表征目标点上的血流速度方向，组合获得沿相应发射角度的血流速度分量，可以采用向量的方式来表达。
在上述实施例中采用的发射角度作为实施例，如果采用前文中提到的沿一个接收角度获得多次超声波信号，则也可以采用上述方式来进行计算，但是前文中的发射角度则替换为接收角度，血流速度方向则为接收角度，从而可以获得沿相应接收角度的血流速度分量。
采用上述多普勒计算方式，根据不同角度的超声波信号，可以分别获得沿不同角度方向上的血流速度值，其可以用多普勒频率来表征。
通常，在超声成像中，利用多普勒原理，对超声波信号进行多普勒处理，可以获得扫描目标或者其内的运动部分的运动速度。例如，获得了超声波信号之后，通过自相关估计方法或者互相关估计方法，可以根据超声波信号获得扫描目标或者其内的运动部分的运动速度。对超声波信号进行多普勒处理以获得扫描目标或者其内的运动部分的运动速度的方法可以使用本领域中目前正在使用或者将来可能使用的任何可以用以通过超声波信号计算扫描目标或者其内的运动部分的运动上速度的方法，在此不再详述。
根据前文的方法获得目标点处沿不同发射角度或接收角度的血流速度分量，然后在目标点处合成这些速度分量，即可获得该目标点处的合成速度，即目标点的血流速度矢量。如图10所示，具体地如下所示。
首先，按照前文所述方式获取至少两组超声波信号，所述至少两组超声波信号可以是分属多个不同角度的超声波信号，不同角度包括不同的发射角度或不同的接收角度(步骤S110)。按照不同角度对应的超声波信号，存储为与角度相关的至少二组数据帧集。步骤S210，按照不同角度对应的超声波信号，存储为与角度相关的至少两组数据帧集。
其次，基于分属不同角度的数据帧集，参照前文利用多普勒成像技术的计算过程，分别计算每一组数据帧集对应的血流速度分量，获得与所述角度相关的至少两个血流速度分量(步骤S220)。在每个目标点处获得至少两个血流速度分量。
然后，将至少两个血流速度分量进行速度合成，获得步骤S200期望获得的血流速度矢量，其中包括合成后获得的血流速度值和合成角度，合成角度指向血流速度方向(步骤S230)。通过此方法可以获得多个目标点处分别对应的血流速度矢量。
在上述实施例中采用的发射角度作为实施例，如果采用前文中提到的沿多个接收角度获得多组超声回波信号，则也可以采用上述方式来进行计算，但是每个步骤中的“发射角度”应被替换为“接收角度”。
当然针对一个发射角度或接收角度对应的血流速度分量，本发明不限于上述方法，还可以采用其他本领域中已知或者未来可能采用的方法。
前文中已提出了多种有关血流速度矢量的计算方式，血流速度矢量中速度值可以包括目标点处的近似或真实的速度、加速度、速度方差评估值等等表征速度状态的统计量中的其中一种。
在步骤S120中或者与步骤120并列的方式还包括以下步骤：
根据上述超声波信号，图像处理模块获得扫描目标的至少一部分的超声图像，本文的超声图像可以是三维超声立体图像，也可以是二维超声图像，例如B图、用以显示的通过上述扫描体获得的三维超声图像数据库中的图像，或者通过二维血流显示技术获得的增强型B图像。在本发明的一个实施例中，超声图像可以使用平面超声波束成像，也可以使用聚焦超声波束成像。但是由于聚焦超声波束每次发射的能力较集中，而且仅在能力集中处成像，因此获得的回波信号信噪比高，获得的超声图像质量较好，而且聚焦超声波束的主瓣狭窄，旁瓣较低，获得的超声图像的横向分辨率也较高。所以，在本发明的一个实施例中，超声图像可以使用聚焦超声波束成像。同时为了获得更加高质量的超声图像，可以在步骤S100 中发射多次发射聚焦超声波束，来实现扫描获得一帧超声图像。
在本发明的一个实施例中，在上述步骤S100中向扫描目标发射多次聚焦超声波束，而在步骤S200中接收聚焦超声波束的回波，获取一组聚焦波束回波信号，根据所述聚焦波束回波信号获得扫描目标的至少一部分的超声图像。利用聚焦超声波可以获得高质量的超声图像。有关平面超声波束和聚焦超声波束的结合发射过程参见前述相关内容。
另外，获取超声图像的数据，可以基于前述步骤S100中任意一组超声波信号或任意一组数据帧集来获得。通过显示器显示超声图像，并还可以在超声图像上显示采样框。本实施例的采样框可以是一个，也可以是多个，并且多个采样框可以重叠。采样多个采样框时，可以同时对比观察多个离散度的量化结果。在其中一个实施例中，采样框可以由用户调节，基于用户对采样框的调节信号，确定重新定义的采样框，获得采样框的大小和形状。
在步骤S300中，图像处理模块7提取多个血流速度方向。
提取的多个血流速度方向至少包括以下内容之一：
1、在同一时间时多个位置处的血流速度方向；和，
2、在同一位置处不同时间对应的多个血流速度方向。
本实施例中，同一时间包括同一时刻，或同一时间段；时刻包括至少一个真实的时间点；时间段包括至少一个时刻。
本实施例中的时刻还可以由图像的采集帧率来确定。采用同一时间多个位置处分别对应的血流速度方向，可以评估此多个位置处在同一时间的涡流等情况，而采样同一位置在不同时间所分别对应的血流速度方向，则可以评估同一个位置处在一定时期内的涡流等情况。例如，在心脏超声心动图中，提取的多个血流速度方向包括同一心动周期中任意时相对应的多个血流速度方向，从而可以评估一个心动周期中同一位置在不同期间或同一期间的血流方向变化。也可以是，提取的多个血流速度方向包括：不同心动周期内同一时相对应的多个血流速度方向，用以量化评估收缩期或舒张期内血流速度方向的变化。本文中的时相包括心动周期中的任意时刻或时间段，包括收缩期和/或舒张期。
本实施例的位置可以为扫描目标内感兴趣的点或者区域，通常表现为，在显示器上展示的扫描目标的至少一部分超声图像中，可被标记或者可被展示的感兴趣的点或区域。例如，在其中一些实施例中，上述位置可以包括至少一个目标点。当一个位置包括多个目标点时，则，该位置处的血流速度方向可以取该位置处多个目标点对应的血流速度矢量的合成角度方向，或者取该位置处多个目标点对应的血流速度方向中的其中之一，或者可以取该位置处多个目标点对应的血流速度 方向中数量最多的一个方向，等等。同理，该位置处的血流速度矢量可以取该位置处的多个目标点对应的血流速度矢量的合成速度，或者取该位置处的多个目标点对应的血流速度矢量中的其中之一，或者取该位置处的多个目标点对应的血流速度矢量中方向数量最多的一个，或者取该位置处的多个目标点对应的血流速度矢量的均值等等。
在其中一些实施例中，根据设定的感兴趣区域来确定待获取血流速度方向的多个位置，此设定的感兴趣区域可以包含用户选择的感兴趣区域，基于系统图像分割技术自动分割获得的脉管区域，系统默认选择的感兴趣区域范围和扫描目标的整个成像区域等其中的一个范围或多个范围的组合。多个位置可以是感兴趣区域的多个离散或连续的位置，可以由系统自动分配，或者由用户选择。可见，步骤S300中，提取感兴趣区域内的多个血流速度方向，用以量化估计感兴趣区域的
在其中一个实施例中，感兴趣区域通过采样框来确定，该采样框可以是系统自动在超声图像上形成的区域，或者是整个成像区域，或者还可以是用户在超声图像上输入选择指令获得区域，等等。通常感兴趣区域包括至少一个目标点，或包含至少一个目标点的领域(数据块)。例如，图13中的31，图14中的41，图15中的51或A32，图16中的61或62，和图17中的71或72。
多个位置，多个时刻，多个时间段都可以是离散提取的，而并非一定是连续的。当然，提取的多个血流速度方向也可以是，上述第1种和第2中的组合，例如，提取的多个血流速度方向包括多个位置分别在多个时间对应的血流速度方向。上述实施例可以从空间和时间维度来综合衡量脉管中血流流动的情况。
此外，在其中一个实施例中，参见基于图5的变形方案图11，图像处理模块执行步骤S600根据超声波信号获得扫描目标的至少一部分的超声图像，显示器执行步骤S700显示超声图像，执行步骤S800利用操作控制模块获取超声图像上的采样框，在步骤S200之后执行步骤S310，图像处理模块提取与采样框关联的多个血流速度方向，多个血流速度方向可以是采样框中离散的多个位置或者连续的多个位置处的血流速度方向；也可以是多个采样框分别对应的多个位置处的血流速度方向。在其中一个实施例中，提取的多个血流速度方向至少包括在同一时间所述采样框内与位置相关的血流速度方向中的至少一部分，和，所述采样框与时间相关的血流速度方向中的至少一部分。提取的多个血流速度方向可以是采样框内的所有位置对应的多个血流速度方向，也可以是部分位置对应的多个血流速度方向。采用此方式时，可以提取采样框中部分期望进行离散化度量的位置，这些期望的位置可能比较接近，也可以距离较远，从而可以更加灵活的宏观上来分 析一定较大范围内的血流速度方向的离散度。
在本实施例中提取多个血流速度方向之前还包括：记录每个位置处的血流速度矢量与时间的对应关系；然后在提取多个血流速度方向时从所述对应关系中提取多个时刻所分别对应的血流速度方向。从而能够更加清晰的获知关于血流速度方向的时间信息，便于后续过程中进行显示。
血流速度方向采用角度值或方向向量表示。例如，由一个0-360度的角度数值来表示血流速度方向，或者由球坐标或直角坐标系下的坐标来表示。
在步骤S400中，图像处理模块量化提取的多个血流速度方向的离散度。离散度表示提取的多个血流速度方向之间的差异程度。量化指的是用值来表示离散度。
本实施例中，量化提取的多个血流速度方向的离散度采用以下方式之一：
计算多个血流速度方向的方差，计算多个血流速度方向的角度差极值，和计算多个血流速度方向的标准差。具体计算方式如下文说明。
参见图7所示，以直角坐标系中向量表达血流速度方向的方式来举例说明方差和标准差的计算方式。
球坐标与直角坐标系转化公式如下公式(7)来表示：
Figure PCTCN2016101208-appb-000015
空间中每个点的速度方向需要用两个角度值才能表示出来，即θ和
Figure PCTCN2016101208-appb-000016
速度大小则用r表示，通过坐标转换，也可以在直角坐标系中用向量的方式表示，即(x，y，z)，它包含了速度的大小和方向信息。多个点上的速度方向(角度)的离散度可以采用下列向量方差或者标准差计算公式。
血流速度矢量的二维向量表达为(x1，y1)，(x2，y2)，......(xN，yN)，通过处理得到单位向量
Figure PCTCN2016101208-appb-000017
等，它们只有血流速度方向信息，(a，b)表示代表血流速度方向的二维单位向量。
Figure PCTCN2016101208-appb-000018
Figure PCTCN2016101208-appb-000019
按照下述公式(9)分别求每个维度的平均值：
Figure PCTCN2016101208-appb-000020
用于量化多个血流速度方向的离散度的二维向量方差Var2可以表示为：
Figure PCTCN2016101208-appb-000021
用于量化多个血流速度方向的离散度的二维向量标准差SD2为：
Figure PCTCN2016101208-appb-000022
血流速度矢量的三维向量：(x1，y1，z1)，(x2，y2，z2)，......(xN，yN，zN)同上方式，通过处理得到单位向量
Figure PCTCN2016101208-appb-000023
等，它们只有方向信息，(a，b，c)表示代表血流速度方向的三维单位向量。
Figure PCTCN2016101208-appb-000024
按照下述公式(13)分别求每个维度的平均值
Figure PCTCN2016101208-appb-000025
Figure PCTCN2016101208-appb-000026
用于量化多个血流速度方向的离散度的三维向量方差Var3可以表示为
Figure PCTCN2016101208-appb-000027
用于量化多个血流速度方向的离散度的标准差SD3为
Figure PCTCN2016101208-appb-000028
二维空间中，多个血流速度方向的角度差极值采用以下方式获得。
在若干个角度数中(假设N个)，计算任意两个角度之间的角度差，找到最大的角度差，即角度差极大值为公式(16)，找到最小的角度差，即角度差极小值为公式(17)。
Figure PCTCN2016101208-appb-000029
Figure PCTCN2016101208-appb-000030
其中，di代表第i个血流速度方向的角度值大小，dj代表第j个血流速度方向的角度值大小。di或dj它们是一个0-360度的角度数值。
无论多少个角度无论什么样的角度，根据以上公式(16)和(17)计算出来的角度差极值是一个0至180度之间变化的数值。数值越大代表角度差极值越大。
采用向量的形式计算二维或者三维空间中多个血流速度方向的角度差极值，参见下述公式(18)和(19)。
二维或者三维空间的计算角度差极大值公式如下：
Figure PCTCN2016101208-appb-000031
Figure PCTCN2016101208-appb-000032
角度差极大值还可以用距离来衡量，例如下述公式(20)和(21)
Figure PCTCN2016101208-appb-000033
Figure PCTCN2016101208-appb-000034
其中，
Figure PCTCN2016101208-appb-000035
等如之前定义，是二维或三维单位向量。得到的是两个向量之间的距离(二维空间或者三维空间都适用)。
可见，在计算多个血流速度方向的角度差极值时，首先计算任意两个角度之间的角度差；然后，查找所述角度差的最大值或最小值。本实施例中的角度差可以如公式(18)和(19)的表达方式，也可以采用公式(20)和(21)中的距离衡量方式。
此外，还可以用另一种方式来计算多个血流速度方向的方差。
血流速度方向量化后是一个0至360度(或者-180至180度)之间的值，可以采用如下公式(22)计算方差Var。
Figure PCTCN2016101208-appb-000036
公式中的θ为血流速度方向量化后的角度值，N代表采样框中目标点的个数或者不同采样框分属的位置个数。Var是一个0-1的数，也可以将其乘以100，则方差是一个0-100的数，依次类推可做相应的变形。Var越大代表在采样框中的速度方向变化越大(离散度越大，存在紊流或者涡流等形态的程度越大)，Var越小则采样框中的速度方向一致性越高，越接近层流，例如当Var为0时，则采样框中所有点的血流速度方向完全一致，为典型的层流状态。
本实施例虽然只提供上述几种方式，但是本发明中并不限制量化多个血流速度方向离散度的方式，还可以采用其他用户选择的方式。当然在系统中还可以提供多种量化方式供用户选择，并进行采用不同量化方式获得的离散度结果的对比查看，实现综合评估。
在步骤S500中，利用显示器显示离散度的量化结果。当然，还可以包括利用显示器显示超声图像，并在超声图像中标记采样框，用于对比查看关联区域处离散度的量化结果。
本实施例中显示离散度的量化结果的方式至少可以采用以下方式之一：
通过文本显示所述离散度的量化结果；
显示图标模型，所述图标模型基于所述量化结果构建；和，
在超声图像上叠加显示离散度的量化结果。
在其中一些实施例中，在超声图像上叠加显示离散度的量化结果的方法包括：
首先，利用图像处理模块产生与特定区域关联的质点区块，所述质点区块的颜色编码与特定区域内血流速度方向的离散度量化结果相关。然后，在所述超声图像的特定区域处显示带颜色编码的质点区块，获得离散度图像。质点区块可参见图13和图14中的色斑和A1，A2，A3，A4的色块分区。
本实施例中的特定区域是指，在超声图像上一个特定区域内的血流速度方向量化结果对应一个质点区块，特定区域可以按照系统设定或用户选定的图像区域范围大小，例如3*8，4*4，5*5。按照特定区域的范围大小分割超声图像，依次计算超声图像中多个特定区域内分别对应的血流速度方向的离散度量化结果，这里的多个特定区域可以是对超声图像进行不重叠分割获得的结果，每个特定区域的面积大小可以相同也可以不相同，例如，超声图像是80*80，可以均分为100个8*8的特定区域，每个特定区域之间不重叠。通过每个特定区域中的各个目标点的血流速度方向，依次计算各个特定区域的血流速度方向的离散度，计算结果就是特定区域内中心点的离散度量化结果。并按照该离散度量化结果的大小对相应特定区域进行颜色编码，显示在超声图像上。
当然，这里的多个特定区域也可以是对超声图像进行连续重叠式分割获得的结果，例如，如图8所示，在整个采样框(参见图8中的大方框)中通过一个小采样框(参见图8中的小方框)逐次偏移一个像素点或多个像素点后移动到下一个位置，获得新的图像区域范围，依次来连续重叠式分割超声图像。小采样框连续偏移获得的多个图像区域范围，对应多个特定区域。在计算图8中整个采样框中的血流速度方向的离散度时，首先选取一个小采样框(参见图8中的小方框)，通过小采样框中各个目标点的血流速度方向计算血流速度方向的离散度，计算结果就是小采样框中心点的离散度量化结果。移动这个小采样框后再次计算得到其中心位置的血流速度方向的离散度。以此类推，得到大采样框中多个重叠的特定区域的血流速度方向的离散度。这个小采样框的大小可变，每次计算后移动或偏移的距离也是可变的。图8中两个相互重叠的方框代表两次计算时的小采样框， 而计算的值就是相应中心点位置上的离散度量化结果，按照该离散度量化结果的大小对相应小采样框依次进行颜色编码，显示在超声图像上形成离散度图像。两个采样框是可以有重叠部分的，重叠越多则计算出来的离散度图像的空间分辨率越高。
图13给出了一种彩色效果的离散度图像，其中在采样框31内的脉管部分叠加显示相应特定区域的离散度量化结果，获得诸如图13中具有多个色块的离散度图像，并且，在界面上提供颜色编码Bar 32，提示用户识别离散度的大小。图14提供的图13的线条效果图，采样框为41，超声图像为42，其中A1，A2，A3，A4分别显示渲染不同的颜色，以示当前区域内离散度的差别。此外，如果采用图8所示的方式来进行渲染，那么A1，A2，A3，A4的色块分区将会进一步细分，使得离散度图像的空间分辨率提高。可见本实施例中提供了一种新的成像效果示意图，能够给医生提供更加直观的观察角度，从与超声图像的叠加显示效果上，展现脉管内的血流运动情况。图13中绿色表示离散度最小，红色表示离散度最大。
在其中一些实施例中，可以基于文本显示的方式，直接在显示界面上的任意一个位置显示离散度的量化结果，例如，图16和图17中显示的“SD 0.01/100”。
在其中一些实施例中，构建基于所述量化结果构建的图标模型，显示图标模型，用以展示关联采样框的离散度量化结果。图标模型的展现方式可以有很多方式，例如，如图16中的642，采用带颜色编码的矩形柱，颜色编码与离散度的量化结果大小相关。还例如，如图16中的641，利用带箭头指向的圆形图标，其中箭头指向与离散度量化结果相关。
还可以，基于位置关系构建图标模型，例如，如图18所示，超声图像内显示区域80中的采样框81，或者说是感兴趣区域91中通过显示表征血流速度矢量的标识83，构建与脉管或脉管的部分图像区域对应的图标模型84，对该图标模，84上进行分区(B1，B2，B3，B4)，计算每个分区对应到超声图像上相应区域内的离散度量化结果，然后将该离散量化结果以文本或者颜色编码的方式叠加在相应分区上，用以在显示器上展现。
此外，还可以基于时间变化构建图标模型，例如，构建离散度与时间的关系坐标，记录离散度的量化结果随时间的变化，显示所述离散度的量化结果随时间的变化，在关系坐标中逐一描绘离散度的量化结果随时间的变化，从而生成与采样框关联的离散度变化图。离散度变化图属于图标模型的一种。参见图15给出了一种显示效果图，图中超声图像50中包括大采样框51，小采样框A32；对应小采样框A32的离散度度量化结果图表53所示，建立离散度(例如方差SD)与 时间的坐标关系，然后提取各个时间t31，t32，t33，t34，t35时的离散度量化结果，将其绘制在坐标关系中，获得图表53所示的离散度变化图，用于展现小采样框A32在不同时间变化时离散度的量化结果。此外在图标53中还标记有黑色倒三角形的标识，提示小采样框A32的位置，还可以用以指示当前时刻对应的位置。此外，图19还提供了另一种显示效果图，图中超声图像90包括大采样框91和小采样框92，对应小采样框92的离散度度量化结果图表93所示，建立时间轴的坐标关系，在时间轴的坐标上绘制各个时间t31，t32，t33，t34，t35等分别对应的离散度量化结果，在时间轴上绘制不同时间对应的矩形柱，每个矩形柱的颜色或高度与当前时间对应的离散度的量化结果相关，形成离散度变化图。
无论采用上述哪种显示方式，离散度的量化结果会与超声图像中的相应位置范围进行关联，当位置范围改变时，则量化结果也会随之更新。例如，通过采样框来标记在超声图像中感兴趣的目标点或目标点区域，那么当获取用户对所述采样框的调节信号，根据所述调节信号，确定重新定义的采样框后，在步骤S300中提取与重新定义的采样框关联的多个血流速度方向，这里与采样框关联的含义包括，在采样框中选择多个血流速度方向，或者还包括，由采样框数量确定提取的血流速度方向的个数。在显示离散度的量化结果中，所述量化结果随所述采样框的更新而改变。这里的改变也包括，量化结果的数量和/或量化结果的值随着采样框的更新而改变。
血流速度方向的离散度量化结果可以单独显示，也可以结合超声图像一起显示在显示器的显示界面上，而这里的超声图像可以是多普勒血流图像、血流抛射图像和B图像等等中的其中之一。也即是在显示离散度的量化结果时，同时显示超声图像，并在超声图像上叠加血流速度矢量。以下提供一种在超声图像上叠加血流速度矢量的显示方式。
在其一个实施例中，上述步骤S200中包括：根据上述步骤S100中获得的超声波信号，计算目标点位于不同时刻的超声图像中第一显示位置处的血流速度矢量，用以获得目标点位于不同时刻的超声图像中的血流速度矢量信息。那么在下述过程中，对比显示的可以是各个时刻超声图像中第一显示位置处的血流速度矢量信息。如图9(a)所示，根据上述步骤S200中获得的超声波信号，可以分别获得t1、t2、……、tn时刻对应的超声图像数据P1、P2、……、Pn中，然后计算目标点在各个时刻超声图像中第一显示位置处(图中黑色圆点的位置)的血流速度矢量。本实施例中，目标点在各个时刻超声图像中第一显示位置始终位于二维图像中的位置(H1、W1)处。基于此，在后续步骤S800中对比显示血流速度矢量信息时，即在显示器显示的超声图像P0中在目标点(H1、W1)处显示不同时 刻对应计算的血流速度矢量。若目标点参照上述具体实施例中根据用户自主选择部分或全部、或者由系统默认，那么对应就可以获知相应的第一显示位置，并通过计算当前时刻对应的超声图像中第一显示位置处的血流速度矢量信息用以对比显示，本文中将这种显示模式称为在超声图像上叠加血流速度矢量的第一模式，下文同。图9(a)实例中给出了二维图像P0显示时的效果示意图，当然也可以应用于三维图像显示中，即将各个时刻的超声图像取为前文提到的扫描体获得三维图像数据库，而第一显示位置取为三维图像数据库中的空间三维立体坐标位置，在此不再累述。
在另一个实施例中，上述步骤S200中包括：根据上述步骤S100中获得的超声波信号，计算目标点连续移动到超声图像中相应位置处而依次获得的血流速度矢量，从而获取目标点的血流速度矢量。在本实施例中，通过重复计算目标点在一时间间隔内从一位置移动到超声图像的另一位置处的血流速度矢量，用以获得目标点从初始位置开始连续移动后在超声图像中各个相应位置处对应的血流速度矢量。也就是说，在本实施例的超声图像中用以确定血流速度矢量的计算位置可以通过计算获得。那么在下述过程中，对比显示的可以是各个时刻超声图像中计算获得的位置处的血流速度矢量。
如图9(b)所示，根据上述步骤S100中获得的超声波信号，可以分别获得t1、t2、......、tn时刻对应的超声图像数据P11、P12、......、P1n中，然后，参照上述实施例中根据用户自主选择目标点的部分或全部、或者由系统默认目标点的密度等，确定目标点的初始位置，如图9(b)中(H1、W1)的第一点，然后计算初始位置在时刻t1超声图像P11中的血流速度矢量A1。其次，计算目标点(即图中黑色圆点)从时刻t1的超声图像P11上的初始位置移动到时刻t2的超声图像P12上的位置(H2、W2)，然后根据超声波信号，获得超声图像P12中(H2、W2)处的血流速度矢量，用以对比显示。比如，沿时刻t1超声图像P11中(H1、W1)上的血流速度矢量的方向，移动一时间间隔(其中，时刻t2-时刻t1＝时间间隔)，计算达到第二时刻t2时的位移，如此在第一个时刻t1上的一个目标点在第二个时刻超声图像上的第二显示位置就找到了，然后再依据上步骤S100中获得的超声波信号获得此第二显示位置上的血流速度矢量，从而得到目标点在时刻t2超声图像P12中血流速度矢量信息。依次类推，每相邻的两个时刻，沿目标点在第一时刻对应的血流速度矢量的方向，移动相邻两个时刻的时间间隔获得位移量，根据位移量确定目标点在第二时刻超声图像上的对应位置，再根据超声波信号获得目标点从第一时刻移动到第二时刻的超声图像中相应位置处的血流速度矢量，依此方式可以获得目标点从超声图像中(H1、W1)处连续移动到(Hn、 Wn)处的血流速度矢量，从而获得目标点从初始位置连续移动到不同时刻的超声图像中相应位置处的血流速度矢量，用以获取目标点的血流速度矢量，使其与超声图像同时显示。
本实施例的显示方式中，计算出目标点在一时间间隔的移动位移、并依据该位移确定超声图像中目标点的相应位置，从初始选择的目标点开始按照该时间间隔移动，这一时间间隔可以由系统发射频率决定，还可以是由显示帧率决定，或者还可以是用户输入的时间间隔，通过按照用户输入的时间间隔计算目标点移动后达到的位置，然后在获得该位置处的血流速度矢量信息用以对比显示。初始时，可以依据前文所述方式在图中标注上N个初始目标点，每个初始目标点上都有箭头来表示这个点流速的大小和方向，如图9(b)所示。在显示的过程中，标记目标点连续移动到相应位置处时对应获得的血流速度矢量，形成随时间呈流动状的标识。通过标记图9(b)方式计算获得的血流速度矢量，那么随时间的变化，在新生成的图中，原来每个点的箭头都会发生位置改变，这样可以用箭头的移动，即可形成类似的血流流动过程，以便用户能观察到近似真实的血流流动显像效果，本文中将这种显示模式称为第二模式，下文同。同样，图9(b)实例中给出了二维图像P10显示时的效果示意图，当然也可以应用于三维图像显示中，即将各个时刻的超声图像取为前文提到的扫描体获得三维图像数据库，而第一显示位置取为三维图像数据库中的空间三维立体坐标位置，在此不再累述。
为了提高显示效果，避免因血流速度显示过快而使人眼无法识别，则在本发明的一个实施例中，上述叠加显示血流速度矢量的过程中还包括在显示血流速度时，对上述步骤S200获得的血流速度矢量进行慢放处理，用以对比显示慢放处理后的血流速度矢量。例如，首先对血流速度矢量进行慢放处理，生成慢速血流速度矢量；然后，在上述超声图像上叠加显示慢速血流速度矢量，形成上述血流抛射图，从而实现血流速度矢量和离散度量化结果的对比显示。
在其中一个实施例中，通过产生质点投射体作为标识描绘目标点上血流速度的变化，质点投射体的颜色编码和/或长度与所述脉管中特定位置处血流速度值相关；并将所述质点投射体送入显示器，在所述超声图像的特定位置处显示所述质点投射体随时间的变化，用以通过质点投射体的动态显示来动态展现所述脉管中血流的运动，从而获得血流抛射图。更进一步的，所述质点投射体还包括方向标识，所述方向标识的指向与所述血流速度的方向相关。利用本实施例的方法可以在显示的血流抛射图中可以清晰的描绘目标点在扫描目标内的实际流向，相比只在图像中的相应显示位置显示当前位置随时间变化的血流速度的大小和方向的方式，可以更加精确、更加真实和形象的表示扫描目标内实际的血流走向。这里 可以通过流动的点或箭头，或者可以描绘方向的其他标志来描述流动的血流流动的过程。参见图15至图19中，利用箭头83来表示质点投射体。
此外，质点投射体也可以仅包含方向标识，而不携带血流速度值信息，方向标识的指向与所述扫描目标中特定位置处的血流速度方向相关。在超声图像的特定位置处显示包含方向标识的质点投射体，用以动态的展现所述扫描目标中血流的运动方向。
本实施例中的质点投射体可以是类似于箭头的表现方式，箭头的长短和/或粗细可以用于表现血流速度值，箭头的指向可以用于表现血流速度方向。本实施例中的特定位置是指，在超声图像上一个特定位置显示的血流速度矢量对应一个质点投射体，特定位置可以是用于标记显示血流速度值的位置，例如可以是图9(a)和图9(b)中提到的第一显示位置或第二显示位置。
用于计算离散度的多个血流速度方向包括：同一心动周期中任意时相(例如收缩期和/或舒张期)对应的多个血流速度方向，或者，不同心动周期内同一时相(例如收缩期或舒张期)对应的多个血流速度方向。当需要对心动周期内的多个时间上的血流速度方向进行选择时可参见图17所示的实施例，超声图像70内包括大采样框71和小采样框72，利用小采样框72选择待查看的位置，同时可以基于显示界面上提供的心电图或者多普勒频谱图74，来通过光标73的选择，定位心脏收缩期和/或舒张期上的不同时间帧对应的血流速度方向，然后计算一个收缩期或舒张期，或一个心动周期内的血流运动方向离散情况。可见，在其中一个实施例中，利用显示器显示表征心动周期的图形，包括心电图、多普勒频谱图、包含心动周期多帧图像的视频浏览轴等等可以直观确定心动周期内任意时间的图形，获取用户在表征心动周期的图形上的选择信号，根据所述选择信号提取同一位置处多个时间分别对应的多个血流速度方向。本实施例中在超声心动图模式下可以非常便利的为用户提供操作指引，供用户自行选择。
图5为一个实施例的参数显示方法的流程示意图。应该理解的是，虽然图5的流程图中的各个步骤按照箭头的指示依次显示，但是这些步骤并不是必然按照箭头指示的顺序依次执行。除非本文中有明确的说明，这些步骤的执行并没有严格的顺序限制，其可以以其他的顺序执行。而且，图5中的至少一部分步骤可以包括多个子步骤或者多个阶段，这些子步骤或者阶段并不必然是在同一时刻执行完成，而是可以在不同的时刻执行，其执行顺序也不必然是依次进行，而是可以与其他步骤或者其他步骤的子步骤或者阶段的至少一部分并行执行或者交替地执行。图10至图12为图5的延伸实施例，相关步骤可参见前文相关说明。
以上各个实施例在具体说明中仅只针对相应步骤的实现方式进行了阐述，然 后在逻辑不相矛盾的情况下，上述各个实施例是可以相互组合的而形成新的技术方案的，而该新的技术方案依然在本具体实施方式的公开范围内。
通过以上的实施方式的描述，本领域的技术人员可以清楚地了解到上述实施例方法可借助软件加必需的通用硬件平台的方式来实现，当然也可以通过硬件，但很多情况下前者是更佳的实施方式。基于这样的理解，本发明的技术方案本质上或者说对现有技术做出贡献的部分可以以软件产品的形式体现出来，该计算机软件产品承载在一个非易失性计算机可读存储载体(如ROM、磁碟、光盘、服务器云空间)中，包括若干指令用以使得一台终端设备(可以是手机，计算机，服务器，或者网络设备等)执行本发明各个实施例所述的方法。
在本实施例中的基于了血流速度矢量的成像方法，首先计算出血流的方向，然后利用这个方向进行进一步评估涡流或者紊流程度，从而作为判断狭窄程度的定量分析方法。通过计算同一时刻不同位置或者同一位置不同时刻血流方向的方差值，得到用来做诊断的具体数据，为医生提供了更加直观的图像分析结果，提升了超声成像系统的智能化。
以上所述实施例仅表达了本发明的几种实施方式，其描述较为具体和详细，但并不能因此而理解为对本发明专利范围的限制。应当指出的是，对于本领域的普通技术人员来说，在不脱离本发明构思的前提下，还可以做出若干变形和改进，这些都属于本发明的保护范围。因此，本发明专利的保护范围应以所附权利要求为准。

Claims (31)

  1. 一种超声血流的参数显示方法，其包括：
    通过探头获得来自于扫描目标内的超声波信号；
    根据所述超声波信号，获得所述扫描目标内的血流速度方向；
    提取多个血流速度方向；
    量化提取的多个血流速度方向的离散度；
    显示所述离散度的量化结果。
  2. 根据权利要求1所述的超声血流的参数显示方法，其中，所述根据所述超声波信号，获得所述扫描目标内的血流速度方向中包括：
    根据所述超声波信号，获得所述扫描目标内的血流速度矢量，所述血流速度矢量包括血流速度值和血流速度方向。
  3. 根据权利要求1或2所述的超声血流的参数显示方法，其中，所述提取多个血流速度方向之前包括：
    根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像，
    显示所述超声图像，
    获取所述超声图像上的采样框；和
    所述提取多个血流速度方向包括：提取与所述采样框关联的多个血流速度方向。
  4. 根据权利要求1所述的超声血流的参数显示方法，其中，所述多个血流速度方向至少包括以下之一：
    在同一时间时多个位置处的血流速度方向；和，
    在同一位置处不同时间对应的多个血流速度方向。
  5. 根据权利要求1所述的超声血流的参数显示方法，其中，所述量化提取的多个血流速度方向的离散度采用以下方式之一：
    计算多个血流速度方向的方差，
    计算多个血流速度方向的角度差极值，和
    计算多个血流速度方向的标准差。
  6. 根据权利要求5所述的超声血流的参数显示方法，其中，所述计算多个血流速度方向的角度差极值包括：
    计算任意两个角度之间的角度差；
    查找所述角度差的最大值或最小值。
  7. 根据权利要求2所述的超声血流的参数显示方法，其中，所述通过探 头获得来自于扫描目标内的超声波信号，根据所述超声波信号，获得所述扫描目标内的血流速度矢量中包括：
    通过探头获得来自于扫描目标内的多个不同角度的超声波信号，所述多个不同角度的超声波信号分属于不同的接收角度或不同的发射角度；
    按照不同角度对应的超声波信号，存储为与角度相关的至少二组数据帧集；
    基于分属不同角度的数据帧集，计算每一组数据帧集对应的血流速度分量，获得与所述角度相关的至少两个血流速度分量；
    合成所述至少两个血流速度分量，获得所述血流速度矢量。
  8. 根据权利要求7所述的超声血流的参数显示方法，其中，所述通过探头获得来自于扫描目标内的超声波信号包括：
    通过探头向扫描目标发射不同发射角度的平面超声波束，
    接收所述平面超声波束的回波，获得分属于不同发射角度的平面超声波信号，用以计算所述血流速度矢量。
  9. 根据权利要求3所述的超声血流的参数显示方法，其中，所述通过探头获得来自于扫描目标内的超声波信号包括：
    通过探头向扫描目标发射聚焦超声波束，
    接收所述聚焦超声波束的回波，获得聚焦超声波信号，用以获得所述超声图像。
  10. 根据权利要求1所述的超声血流的参数显示方法，其中，所述显示所述离散度的量化结果至少包括以下方式之一：
    通过文本显示所述离散度的量化结果；和，
    显示图标模型，所述图标模型基于所述量化结果构建。
  11. 根据权利要求1所述的超声血流的参数显示方法，其中，所述显示所述离散度的量化结果包括：
    根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；
    显示所述超声图像；
    产生质点区块，所述质点区块的颜色编码与特定区域内血流速度方向的离散度量化结果相关；
    在所述超声图像的特定区域处显示带颜色编码的质点区块。
  12. 根据权利要求1所述的超声血流的参数显示方法，其中，所述方法中根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；
    显示所述超声图像时，产生质点投射体，所述质点投射体的颜色编码和/或长度与特定位置处的血流速度值相关；
    在超声图像的特定位置处显示所述质点投射体，用以动态的展现所述扫描目标中血流的运动。
  13. 根据权利要求1所述的超声血流的参数显示方法，其中，所述方法中根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；
    显示所述超声图像时，产生包含方向标识的质点投射体，所述方向标识的指向与所述扫描目标中特定位置处的血流速度方向相关；
    在超声图像的特定位置处显示所述质点投射体，用以动态的展现所述扫描目标中血流的运动方向。
  14. 根据权利要求11所述的超声血流的参数显示方法，其中，相邻的所述特定区域之间重叠。
  15. 根据权利要求3所述的超声血流的参数显示方法，其中，所述提取与所述采样框关联的多个血流速度方向中，提取的多个血流速度方向至少包括以下之一：
    在同一时间所述采样框内与位置相关的血流速度方向中的至少一部分，和，
    所述采样框与时间相关的血流速度方向中的至少一部分。
  16. 根据权利要求3所述的超声血流的参数显示方法，其中，所述方法包括：
    获取用户对所述采样框的调节信号；
    根据所述调节信号，确定重新定义的采样框；
    所述提取与所述采样框关联的多个血流速度方向中，提取与重新定义的采样框关联的多个血流速度方向；
    在所述显示所述离散度的量化结果中，所述量化结果随所述采样框的更新而改变。
  17. 根据权利要求2所述的超声血流的参数显示方法，其中，所述提取多个血流速度方向之前包括：
    记录每个位置处的血流速度矢量与时间的对应关系；和
    所述提取多个血流速度方向包括：
    从所述对应关系中提取多个时间所分别对应的血流速度方向。
  18. 根据权利要求3所述的超声血流的参数显示方法，其中，所述采样 框包括一个或多个。
  19. 根据权利要求1所述的超声血流的参数显示方法，其中，所述显示所述离散度的量化结果之前包括：
    记录所述离散度的量化结果随时间的变化；
    所述显示所述离散度的量化结果包括：
    显示所述离散度的量化结果随时间的变化，生成与采样框关联的离散度变化图。
  20. 根据权利要求1所述的超声血流的参数显示方法，其中，所述提取的多个血流速度方向包括：
    同一心动周期中任意时相对应的多个血流速度方向，或者，
    不同心动周期内同一时相对应的多个血流速度方向。
  21. 一种超声成像系统，其包括：
    探头，用于向扫描目标发射超声波束；
    接收电路和波束合成模块，用于接收来所述超声波束的回波信号，进行波束合成后获得超声波信号；
    图像处理模块，用于根据所述超声波信号，获得所述扫描目标内的血流速度方向，提取多个血流速度方向，量化提取的多个血流速度方向的离散度；
    显示器，用于显示所述离散度的量化结果。
  22. 根据权利要求21所述的超声成像系统，其中，所述图像处理模块，还用于根据所述超声波信号，获得所述扫描目标内的血流速度矢量，所述血流速度矢量包括血流速度值和血流速度方向。
  23. 根据权利要求21或22所述的超声成像系统，其中，所述图像处理模块，还用于根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；获取所述超声图像上的采样框；提取与所述采样框关联的多个血流速度方向；
    显示器，还用于显示所述超声图像。
  24. 根据权利要求21所述的超声成像系统，其中，所述多个血流速度方向至少包括以下之一：
    在同一时间时多个位置处的血流速度方向；和，
    在同一位置处不同时间对应的多个血流速度方向。
  25. 根据权利要求22所述的超声成像系统，其中，所述探头还用于向扫 描目标发射不同发射角度的平面超声波束；所述接收电路和波束合成模块，还用于接收所述平面超声波束的回波，获得分属于不同发射角度的所述超声波信号，图像处理模块基于分属不同角度的超声波信号，计算与所述角度相关的至少两个血流速度分量，合成所述至少两个血流速度分量，获得所述血流速度矢量；
    所述探头还用于通过探头向扫描目标发射聚焦超声波束；所述接收电路和波束合成模块，还用于接收所述聚焦超声波束的回波，获得聚焦超声波信号，图像处理模块根据聚焦超声波信号获得所述超声图像。
  26. 根据权利要求21所述的超声成像系统，其中，所述显示器上至少采用以下方式显示所述离散度的量化结果：
    通过文本显示所述离散度的量化结果；和
    显示图标模型，所述图标模型基于所述量化结果构建。
  27. 根据权利要求21所述的超声成像系统，其中，所述图像处理模块用于根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；
    利用所述显示器显示所述超声图像；
    所述图像处理模块产生质点区块，所述质点区块的颜色编码与特定区域内血流速度方向的离散度量化结果相关；并利用显示器在所述超声图像的特定区域处显示带颜色编码的质点区块。
  28. 根据权利要求21所述的超声成像系统，其中，所述图像处理模块用于根据所述超声波信号，获得所述扫描目标的至少一部分的超声图像；利用所述显示器显示所述超声图像；所述图像处理模块产生质点投射体，所述质点投射体的颜色编码和/或长度与特定位置处的血流速度值相关；并利用显示器在超声图像的特定位置处显示所述质点投射体，用以动态的展现所述扫描目标中血流的运动。
  29. 根据权利要求23所述的超声成像系统，其中，在显示器的显示界面上显示所述离散度的量化结果随时间的变化，生成与采样框关联的离散度变化图。
  30. 根据权利要求21所述的超声成像系统，其中，所述提取的多个血流速度方向包括：
    同一心动周期中任意时相对应的多个血流速度方向，或者，
    不同心动周期内同一时相对应的多个血流速度方向。
  31. 根据权利要求21所述的超声成像系统，其中，所述系统包括操作控 制模块，
    利用显示器显示表征心动周期的图形，利用操作控制模块获取用户在表征心动周期的图形上的选择信号，图像处理模块根据所述选择信号提取多个时间分别对应的多个血流速度方向，用以量化提取的多个血流速度方向的离散度。
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