WO2017035838A1 - 超声灰阶成像系统及方法 - Google Patents

超声灰阶成像系统及方法 Download PDF

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Publication number
WO2017035838A1
WO2017035838A1 PCT/CN2015/088985 CN2015088985W WO2017035838A1 WO 2017035838 A1 WO2017035838 A1 WO 2017035838A1 CN 2015088985 W CN2015088985 W CN 2015088985W WO 2017035838 A1 WO2017035838 A1 WO 2017035838A1
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fluid
image
unfocused
ultrasound
data
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PCT/CN2015/088985
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English (en)
French (fr)
Inventor
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 PCT/CN2015/088985 priority Critical patent/WO2017035838A1/zh
Priority to CN201580009369.1A priority patent/CN106102588B/zh
Publication of WO2017035838A1 publication Critical patent/WO2017035838A1/zh
Priority to US15/912,182 priority patent/US20180192988A1/en
Priority to US18/141,691 priority patent/US12004898B2/en

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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01S—RADIO DIRECTION-FINDING; RADIO NAVIGATION; DETERMINING DISTANCE OR VELOCITY BY USE OF RADIO WAVES; LOCATING OR PRESENCE-DETECTING BY USE OF THE REFLECTION OR RERADIATION OF RADIO WAVES; ANALOGOUS ARRANGEMENTS USING OTHER WAVES
    • G01S15/00—Systems using the reflection or reradiation of acoustic waves, e.g. sonar systems
    • G01S15/88—Sonar systems specially adapted for specific applications
    • G01S15/89—Sonar systems specially adapted for specific applications for mapping or imaging
    • G01S15/8906—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques
    • G01S15/8979—Combined Doppler and pulse-echo imaging systems
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • 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
    • A61B8/46—Ultrasonic, sonic or infrasonic diagnostic devices with special arrangements for interfacing with the operator or the patient
    • A61B8/461—Displaying means of special interest
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/52—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/5215—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of medical diagnostic data
    • A61B8/5238—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of medical diagnostic data for combining image data of patient, e.g. merging several images from different acquisition modes into one image
    • A61B8/5246—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of medical diagnostic data for combining image data of patient, e.g. merging several images from different acquisition modes into one image combining images from the same or different imaging techniques, e.g. color Doppler and B-mode
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01S—RADIO DIRECTION-FINDING; RADIO NAVIGATION; DETERMINING DISTANCE OR VELOCITY BY USE OF RADIO WAVES; LOCATING OR PRESENCE-DETECTING BY USE OF THE REFLECTION OR RERADIATION OF RADIO WAVES; ANALOGOUS ARRANGEMENTS USING OTHER WAVES
    • G01S15/00—Systems using the reflection or reradiation of acoustic waves, e.g. sonar systems
    • G01S15/88—Sonar systems specially adapted for specific applications
    • G01S15/89—Sonar systems specially adapted for specific applications for mapping or imaging
    • G01S15/8906—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/54—Control of apparatus or devices for radiation diagnosis
    • A61B6/547—Control of apparatus or devices for radiation diagnosis involving tracking of position of the device or parts of the device
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/52—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/5207—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of raw data to produce diagnostic data, e.g. for generating an image
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01S—RADIO DIRECTION-FINDING; RADIO NAVIGATION; DETERMINING DISTANCE OR VELOCITY BY USE OF RADIO WAVES; LOCATING OR PRESENCE-DETECTING BY USE OF THE REFLECTION OR RERADIATION OF RADIO WAVES; ANALOGOUS ARRANGEMENTS USING OTHER WAVES
    • G01S15/00—Systems using the reflection or reradiation of acoustic waves, e.g. sonar systems
    • G01S15/88—Sonar systems specially adapted for specific applications
    • G01S15/89—Sonar systems specially adapted for specific applications for mapping or imaging
    • G01S15/8906—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques
    • G01S15/8909—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques using a static transducer configuration
    • G01S15/8915—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques using a static transducer configuration using a transducer array
    • G01S15/8927—Short-range imaging systems; Acoustic microscope systems using pulse-echo techniques using a static transducer configuration using a transducer array using simultaneously or sequentially two or more subarrays or subapertures

Definitions

  • the present invention relates to ultrasound imaging systems, and more particularly to an ultrasonic blood flow imaging method that directly displays blood flow using a gray scale map.
  • Gray-scale blood flow imaging is an ultrasonic blood flow imaging method in which the obtained ordinary B-picture signal is processed and the gray-scale image is used to directly display the blood flow.
  • Traditional gray-scale blood flow imaging uses focused wave emission and line-by-line scanning, which results in higher blood flow accuracy and sensitivity, enhances intravascular flow signals through wall filtering, and then combines tissue signals to connect vessel walls and The flow of red blood cells in the blood vessels is simultaneously displayed in gray scale images.
  • the dynamic map of gray-scale blood flow imaging can visually show the difference between flowing blood and non-flowing tissue.
  • the conventional method uses focused wave emission and line-by-line scanning, which reduces the temporal resolution of the image, and the maximum features of gray-scale blood flow imaging cannot be fully utilized. When the blood flow rate is too fast, there is a scan time interval between each scan line, which will distort the display of gray-scale blood flow imaging.
  • an ultrasonic gray scale imaging system comprising:
  • a transmitting circuit for exciting the probe to emit a plurality of sets of unfocused ultrasonic beams to the fluid
  • a receiving circuit and a beam combining module configured to receive echoes of the plurality of sets of unfocused ultrasonic beams, Obtaining multiple sets of unfocused ultrasound echo signals
  • a signal processing module configured to perform signal detection and enhancement processing on the plurality of sets of unfocused ultrasonic echo signals, and obtain fluid display data
  • An image processing module configured to perform data conversion on the fluid display data to obtain a B-mode ultrasound image sequence
  • an ultrasonic gray scale imaging system comprising:
  • a transmitting circuit for exciting the probe to emit a plurality of sets of unfocused ultrasonic beams and focused ultrasonic beams to a scanning target containing a fluid
  • a receiving circuit and a beam combining module configured to receive echoes of the plurality of sets of unfocused ultrasonic beams returning from the fluid, obtain a plurality of sets of unfocused ultrasonic echo signals, and receive the plurality of sets of focused ultrasonic beams from the scan The echoes returned on the target obtain multiple sets of focused ultrasound echo signals;
  • a signal processing module configured to perform signal detection and enhancement processing on the plurality of sets of unfocused ultrasonic echo signals, and obtain fluid display data
  • a B-mode signal processing module configured to perform ultrasonic signal data after performing signal detection and enhancement processing on the plurality of sets of focused ultrasonic echo signals
  • An image processing module configured to superimpose the fluid display data on the ultrasound image data, and perform data conversion to obtain a B mode ultrasound image sequence
  • an ultrasonic grayscale imaging method comprising:
  • the B-mode ultrasound image sequence is displayed to present a visual effect of fluid dynamic flow.
  • the invention adopts unfocused wave (such as plane wave or divergent wave) emission to obtain multiple scan lines or even the entire image at a time, which can greatly improve the time separation rate of the ultrasonic image, and solve the traditional gray-scale blood flow imaging in displaying high-speed blood.
  • unfocused wave such as plane wave or divergent wave
  • FIG. 1a, 1b, 1c, and 1d are block diagrams showing an ultrasonic gray scale imaging system according to four embodiments 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 a focused ultrasonic beam according to an embodiment of the present invention.
  • Figure 5 is a schematic view showing a diverging ultrasonic beam in an embodiment of the present invention.
  • FIG. 6 is a schematic flow chart of an embodiment of the present invention.
  • FIG. 7 is a schematic flowchart of another embodiment of the present invention.
  • Figure 8 is a schematic diagram of a multi-angle emission planar ultrasonic beam in an embodiment of the present invention.
  • Figure 9 is a schematic illustration of a multi-angle emission unfocused ultrasound beam in one embodiment of the present invention.
  • the ultrasonic gray scale imaging system comprises: 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. .
  • the above signal processing module 6 and image processing module 7 can be implemented based on one or more processors.
  • 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 the transmitted pulse, and transmits an ultrasonic beam to the scanning target containing the fluid, and receives it after a certain delay.
  • the scanning target includes blood vessels existing in organs, tissues, and the like in the human body or animal body or other blood vessels in the living body having fluid flow therein, which are not shown in the drawings.
  • Fluids include fluids present in tissues such as blood vessels, lymphatic systems, etc. in the human or animal body, or flowing fluids present in other vessels in the human or animal body, not shown.
  • the receiving circuit receives the electrical signals generated by the conversion of the probe 1 to obtain ultrasonic echo signals, and sends the ultrasonic echo signals to the beam combining module 5.
  • the beam synthesizing module 5 performs signal processing such as focus delay, weighting, and channel summation on the ultrasonic echo signal, and then sends the ultrasonic echo signal to the signal processing module 6 for signal detection, enhancement processing, and the like, to obtain a fluid containing the same.
  • the fluid of the information shows the data.
  • the image processing module 7 performs different data conversion on the fluid display data according to different imaging modes required by the user, obtains image data of different modes, and then forms different modes of B through logarithmic compression, dynamic range adjustment, digital scan conversion, and the like.
  • the mode ultrasound image while obtaining a series of ultrasound images in the order of the ultrasound acquisition time, is referred to herein as an ultrasound image sequence.
  • the image processing module 7 sends the generated sequence of ultrasound images to the display 8 for display, and a dynamic image can be obtained, thereby exhibiting a visual effect of fluid dynamic flow.
  • Probe 1 typically includes an array of multiple array elements. Each time the ultrasound is transmitted, all 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. At this time, 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, and the formation is transmitted to The synthetic ultrasonic beam of the fluid, the direction of which is the emission direction 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.
  • the propagation direction of the above-described synthetic ultrasonic beam can be changed by controlling the delay between the time at which the element participating in the transmission of the ultrasonic wave is excited by the transmitted pulse.
  • the superposition mentioned in this article can be added to the normal of the two, and can also be added according to a certain weight.
  • 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, "focusing" the ultrasonic waves emitted by each array element to the predetermined position Positioned, the predetermined position of the focus is referred to as the "focus” such that the resulting synthesized ultrasonic beam is the beam focused at that focus, referred to herein as the "focused ultrasound beam.”
  • Figure 4 is a schematic diagram of a focused focused ultrasound beam.
  • the array elements participating in the transmission of the ultrasonic waves in FIG.
  • the ultrasonic waves emitted by each element are focused at the focus to form a focused ultrasound beam.
  • the ultrasonic waves emitted by the respective array elements participating in the transmission of the ultrasonic waves may not be focused during the propagation, and may not be completely diverged. Instead, a plane wave that is generally planar as a whole is formed. In this paper, this non-focal plane wave is called a "planar ultrasonic 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.”
  • the plurality of array elements arranged linearly are simultaneously excited by the electric pulse signal, and each array element simultaneously emits ultrasonic waves, and the emission direction of the synthesized ultrasonic beam is consistent with the normal direction of the array plane of the array elements. For example, as shown in FIG.
  • the plane wave of the vertical emission at this time, there is no time delay between the respective 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 emission pulse), and each array element is
  • the transmitting pulse is simultaneously excited, and 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 arrangement plane of the array element.
  • the angle between the normal directions is zero degrees.
  • 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 deflection angle of the combined beam, changing the above time delay, the magnitude of the deflection angle of the combined beam and the deflection in the normal direction of the array plane of the array element can be adjusted.
  • Figure 3 shows a plane wave that is deflected and emitted. At this time, there is a predetermined delay between each element participating in the transmission of the ultrasonic wave (ie, each element is excited by the transmitted 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 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 deflection of the ultrasonic beam of the plane.
  • the delay time By changing the delay time, the size of the angle a can be adjusted.
  • the diverging ultrasonic beam can be regarded as a non-focusing wave whose virtual focus is behind the probe, and therefore the delay between the time when the array element participating in the transmission of the ultrasonic wave is excited by the transmitting pulse is adjusted.
  • the position of the virtual focus can be adjusted to change the direction in which the unfocused beam is emitted.
  • the direction and the element of the combined beam can be adjusted by adjusting the delay between the time when the array element participating in the transmission of the ultrasonic wave is excited by the transmitted pulse.
  • the combined beam may be the planar ultrasonic beam, the focused ultrasonic beam or the above-mentioned
  • the ultrasonic beam and the like are diverged, and the planar ultrasonic beam and the divergent ultrasonic beam are collectively referred to herein as a non-focused ultrasonic beam.
  • the present invention employs a non-focused ultrasound beam emission scheme that can be used to obtain multiple scan lines or to obtain a single image.
  • Wall filtering of the collected signals can obtain blood flow signals.
  • each line needs to be scanned several times and needs to overcome the transient problem of the filter, which will be more conducive to blood flow signals.
  • Acquisition which can greatly improve the time separation rate of ultrasound images, and solve the distortion problem of traditional gray-scale blood flow imaging in displaying high-speed blood flow.
  • the planar ultrasonic beam and the divergent ultrasonic beam are collectively referred to herein as a non-focused ultrasonic beam.
  • Unfocused ultrasound beams such as each image obtained from a plane ultrasonic beam, are continuously obtained.
  • the transients are not considered in the filtering, and many images can be used for wall filtering to improve the signal-to-noise ratio, and the frame rate is not lost. .
  • Figure 6 shows an ultrasound imaging method.
  • step S100 the transmitting circuit 2 excites the probe 1 to emit a plurality of sets of unfocused ultrasonic beams to the fluid.
  • the unfocused ultrasonic beam here may be a planar ultrasonic beam or a divergent ultrasonic beam.
  • the plurality of sets of unfocused ultrasound beams may be unfocused ultrasound beams that are excited in chronological order. Receiving echoes of the plurality of unfocused ultrasonic beams returning from the fluid, obtaining a plurality of sets of ultrasonic echo signals for forming an image sequence having a certain time series, thereby obtaining source data of the dynamic image display.
  • the receiving circuit 4 and the beam combining module 5 receive a set of unfocused ultrasonic beams, and a set of unfocused ultrasonic echo signals can be obtained.
  • each set of unfocused ultrasonic beams emitted to the fluid can be The non-focusing ultrasonic beam including a plurality of different emission angles
  • the receiving circuit 4 and the beam combining module 5 receive echoes of a plurality of unfocused ultrasonic beams of different emission angles, and obtain a plurality of non-focusings included in a set of unfocused ultrasonic echo signals
  • the ultrasonic echo signal is spatially combined according to the plurality of unfocused ultrasonic echo signals for the non-focused ultrasonic echo signals obtained by the reflection at the same spatial position, and then sent to the signal processing module 6.
  • An echo of an unfocused ultrasonic beam based on a plurality of different emission angles for synthesizing a pair of ultrasound images Therefore, on the basis of ensuring the imaging frame rate, the obtained echo signal has a higher signal-to-noise ratio, and can be used to obtain better quality ultrasonic image data.
  • the probe is excited by a wide bandwidth of encoded pulses to emit an unfocused ultrasound beam to the fluid.
  • the echo signal can be enhanced by using a wide bandwidth coded pulse.
  • the bandwidth mentioned here refers to the frequency range occupied by various frequency components contained in the signal, and the wide bandwidth refers to the frequency range of 3 megabytes to 10 megabytes or even more.
  • each element of the probe uses a coding sequence of the base sequence as a drive pulse, and each pulse in the sequence is commonly referred to as a slice (CHIP).
  • the base sequence is phase encoded using an N-bit transmit code to produce N-coded sequences that are stored in a transmit sequence memory (not shown).
  • Each code sequence read from the transmit sequence memory controls the activation of the transmit circuit 2 during the respective transmit firing.
  • each set of unfocused ultrasound beams that the transmitting circuit 2 excites the probe 1 to emit into the fluid may include multiple times of emission to the fluid, and each shot corresponds to one unfocused ultrasound echo. A signal that is used to obtain a better display of ultrasound images through spatial recombination.
  • the unfocused ultrasound beams that are emitted multiple times to the fluid may have the same emission angle or different emission angles.
  • Non-focusing of multiple different emission angles may be included in each set of unfocused ultrasound beams emitted to the fluid
  • an ultrasonic beam or a plurality of unfocused ultrasonic beams respectively including a plurality of different emission angles alternately emit non-focused ultrasonic beams to the fluid according to different emission angles, or alternately emit multiple unfocused ultrasonic waves to the fluid according to different emission angles. bundle.
  • the probe 1 emits a plane ultrasonic beam to the fluid at a plurality of angles, and different line patterns are used to distinguish different emission angles.
  • the positional ultrasonic echo signals of the oblique line regions in FIG. 8 will be superimposed.
  • each set of unfocused ultrasound beams that the transmitting circuit 2 excites the probe 1 to emit to the fluid may also include unfocused ultrasound beams at a plurality of different positions of the virtual focus.
  • the unfocused ultrasound beam is alternately emitted toward the fluid according to the position of the virtual focus, or the non-focused ultrasound beam is alternately emitted to the fluid in accordance with the position of the virtual focus.
  • the virtual focus positions are sequentially from A, B, and C, and the corresponding unfocused ultrasonic beams are respectively indicated by broken lines, solid lines, and dotted lines.
  • the probe 1 is excited N times by the supplied pulse train to emit N times of the same unfocused ultrasonic beam to the fluid.
  • the transmitting circuit 2 drives the probe so that the generated ultrasonic energy is guided or manipulated to cover the entire scanning surface in a row of ultrasonic beams, which is faster in frame rate than the conventional gray-scale blood flow imaging system adopts focusing super-wave.
  • the acquisition speed of the image data is better, and the non-focused ultrasound beam can be emitted to the desired deflection angle or virtual focus position by appropriately adjusting the time delay of the emission focus.
  • step S200 the receiving circuit 4 and the beam combining module 5 receive the echoes of the plurality of sets of unfocused ultrasonic beams emitted in the above step S100 to obtain a plurality of sets of unfocused ultrasonic echo signals.
  • the ultrasonic echo signals generated by each of the ultrasonic energy bursts are reflected from objects that are located in a series of ranges along each of the transmit beams.
  • the ultrasonic echo signals are detected by the probe 1 respectively, and the sampling of the amplitude of the ultrasonic echo signal at a certain point in time represents the amount of reflection occurring at a particular range. Due to the difference in propagation paths between the reflection point and each element in the probe 1, the ultrasonic echo signals are not detected simultaneously and their amplitudes are not equal.
  • the receiving circuit 4 receives the electrical signals generated by the conversion of the probe 1 to obtain corresponding ultrasonic echo signals, and sends the ultrasonic echo signals to the beam combining module 5.
  • the beam combining module 5 performs focus delay, weighting and pass on the ultrasonic echo signals Tao Qiu and other processing.
  • the beam synthesis module 5 tracks the direction of the transmitted ultrasonic beam and samples the ultrasonic echo signals along a series of ranges of each ultrasonic beam.
  • the beam synthesis module 5 assigns an appropriate time delay and received apodization weight to each of the ultrasonic echo signals, and sums the signals to obtain a synthesized ultrasonic echo signal, which accurately represents the along The sum of the corresponding signals of the plurality of receiving channels within a certain range of the unfocused ultrasonic beam.
  • Beamforming can be achieved by setting the receive beam summer. For the emission of the ultrasonic beam with different emission angles in the same period, the echo signal is obtained by the receiving circuit.
  • the receiving channel corresponding to each array element in the receiving circuit includes an analog-to-digital converter (not shown).
  • the receive beamformer memory assigns an appropriate receive focus time delay to each received echo signal and sums the echo signals to obtain a composite echo signal that accurately represents the reflected from each scan position. Total ultrasound energy. For each scan position, the time delayed received signal is summed at the receive beam summer.
  • the transmitting circuit 2 when the transmitting circuit 2 excites the ultrasonic beam of a plurality of different emission angles in each of the sets of unfocused ultrasonic beams emitted by the probe 1 to the fluid, correspondingly receiving the ultrasonic beams of the plurality of different emission angles
  • the echo obtains a plurality of unfocused ultrasonic echo signals.
  • spatial recombination of the unfocused ultrasonic echo signals obtained by reflection at the same spatial position may be after beam synthesis.
  • step S300 the signal processing module 6 performs signal detection and enhancement processing on the unfocused ultrasonic echo signal to obtain fluid display data.
  • the signal processing here may also include logarithmic compression and the like.
  • Signal detection can include signal envelope detection processing. Unexplained here, see the common methods of signal processing.
  • the ultrasound imaging system shown in FIG. 1b further includes a wall filter 10.
  • the wall filter 10 performs wall filtering processing on the plurality of sets of unfocused ultrasonic echo signals to obtain a filtered signal, and the filtered signal is sent to the signal processing module 7 for signal detection, enhancement processing, and the like to obtain wall filtering.
  • the fluid after the display shows the data.
  • the wall filter 10 may employ a conventional FIR, or IIR, or more complex such as a linear regression filter, a low rank filter, or the like.
  • the value of each of the same positions in the unfocused ultrasonic echo signals at different times is filtered by a high-pass filter, that is, The high frequency blood flow signal is retained, and the low frequency non-fluid signal is filtered out.
  • the echo signal after the wall filtering process will have a higher signal to noise ratio, and obtain better image quality of the ultrasound image data.
  • each of the sets of unfocused ultrasonic beams emitted to the fluid includes a plurality of unfocused ultrasonic beams of different emission angles, and the receiving circuit 4 and the beam combining module 5 receive the plurality of unfocused ultrasonic waves of different emission angles.
  • a plurality of unfocused ultrasound echo signals in a set of unfocused ultrasound echo signals; an unfocused ultrasound echo obtained for reflection at the same spatial position based on the plurality of unfocused ultrasound echo signals The signals are spatially combined and sent to the wall filter 10.
  • the wall filter 10 includes two portions, a first portion for extracting a fundamental frequency component and a second portion for suppressing a fundamental wave component using a high-pass filter.
  • the multiple unfocused ultrasonic echo signals can be subjected to wall filtering processing by one or more wall filters, respectively.
  • step S400 the image processing module 7 performs data conversion on the fluid display data to obtain a B-mode ultrasound image sequence.
  • the image processing module 7 performs data conversion on the fluid display data or the wall-filtered fluid display data after correlation signal processing (including edge enhancement and logarithmic compression, etc.) to obtain an ultrasound image displayed on the display 8. sequence.
  • correlation signal processing including edge enhancement and logarithmic compression, etc.
  • the quadrature demodulated signals after the relevant signal processing include two paths of I and Q, and the two paths of I and Q can be separately subjected to wall filtering processing.
  • the corresponding image data can be obtained in the following manner.
  • the image processing module 7 uses the image processing module 7 to calculate the display variance for the I and Q demodulated data in the fluid display data, to obtain the non-fluid and fluid image regions according to the display variance, and to display the variance by the image gray scale and/or the image color information.
  • the trend is changed to obtain a B-mode ultrasound image sequence superimposed with image gradation and/or image color information, or to obtain a B-mode ultrasound image sequence in which the image gradation and/or image color information is superimposed.
  • a large variance indicates a fluid, and a small variance indicates a non-fluid that is stationary.
  • the formula for calculating the variance can be as follows:
  • the display variance Var is expressed as the following formula (1):
  • I i I-channel demodulated data at the i-th time
  • Q i Q-channel demodulated data at the i-th time.
  • the display variance Var can also be expressed as the following formula (2).
  • the display variance at the corresponding display position calculated by the above formula (1) and formula (2) different image gradations and/or image color information are mapped according to the change trend of the display variance, but obtained by the image processing module 7
  • the image gradation and/or image color information is superimposed in the B mode ultrasound image to achieve a display effect.
  • the overall image gradation information is mapped to the display variance of each display position in the entire pixel image, and then the B-mode ultrasound image sequence obtained by superimposing the image gradation in the B-mode ultrasound image is displayed after continuous playback.
  • the full-frame image exhibits a grayscale effect and exhibits a dynamic cloud-like cluster tumbling visual effect in the fluid flow region.
  • the overall image color information is mapped to the display variance of each display position in the entire pixel image, and then the super-image color information is superimposed on the B-mode ultrasound image, and the obtained B-mode ultrasound image sequence is displayed after continuous playback.
  • the full-frame image exhibits a color contrast effect, and presents a dynamic color cloud-like cluster tumbling visual effect in the fluid flow region.
  • the trend of the magnitude of the variance maps the color of the image at different display positions, and the corresponding image color information can be obtained.
  • the data conversion or the B mode ultrasound image sequence may also be performed in one of the following ways.
  • the I and Q demodulated data in the wall-filtered fluid display data are converted into a polar coordinate system and a Cartesian coordinate system, and the signal envelope is mapped by image gradation and/or image color information.
  • the time-varying amplitude trend is obtained to obtain the above B-mode ultrasound image sequence.
  • directly display the blood flow image Flow_image(x,z), and the conversion mode of the B-mode ultrasound image sequence is as follows:
  • the leg-zero method is used to display the energy of the blood flow signal, and the calculation method can be referred to the above formula (3).
  • the Hilbert transform can be performed on the RF data of the image along the depth (Depth) to obtain the I and Q data, and then the wall filtering process is performed, and the B mode is obtained from one of the above three methods.
  • Ultrasound image sequence When the image data has only one channel of data, the Hilbert transform can be performed on the RF data of the image along the depth (Depth) to obtain the I and Q data, and then the wall filtering process is performed, and the B mode is obtained from one of the above three methods.
  • the image processing module is further configured to calculate signal energy of the characterized fluid according to the wall-filtered fluid display data, and divide the fluid position area and the non-fluid position area in the ultrasonic image data based on the energy threshold, and superimpose the fluid position area on the above
  • the wall-filtered fluid displays data, and the B-mode ultrasound image sequence is obtained after data conversion. This allows the obtained image data to obtain clearer and more accurate fluid image information.
  • different image gradations and/or image color information are mapped according to the magnitude of the signal energy, but in the B-mode ultrasound image obtained by the image processing module 7.
  • the signal energy of each display position in the entire pixel image is mapped to the overall image gradation information, and then the dynamic image vision of the B-mode ultrasound image sequence obtained after superimposing the image gradation in the B-mode ultrasound image is displayed after continuous playback.
  • the full-frame image exhibits a grayscale effect and exhibits a dynamic cloud-like cluster tumbling visual effect in the fluid flow region.
  • the signal color of each display position in the entire pixel image is mapped to the overall image color information, and then the superimposed image color information in the B mode ultrasound image is obtained, and the obtained dynamic image of the B mode ultrasound image sequence is displayed after continuous playing.
  • the full-frame image exhibits a color contrast effect, and presents a dynamic color cloud-like cluster tumbling visual effect in the fluid flow region.
  • step S500 display 8 displays a sequence of B-mode ultrasound images to present a visual effect of fluid dynamic flow. For example, a visual effect of a dynamic gray or colored cloud-like cluster tumbling in the fluid flow region is obtained at the interface of the display 8.
  • the ultrasonic imaging system shown in FIG. 1c further includes a B-mode signal processing module 11, wherein the plurality of sets of unfocused ultrasonic echo signals are output in two ways, and one path is sequentially input to the wall filter. 10.
  • the signal processing module 6 obtains the wall-filtered fluid display data, and the other input to the B-mode signal processing module 11 for signal detection and enhancement processing, the ultrasonic image data is obtained; the image processing module 7 is further used to The wall-filtered fluid display data is superimposed on the above-mentioned ultrasonic image data, and the data conversion is performed to obtain the B-mode ultrasonic image sequence.
  • the signal process of the B mode signal processing module 11 can be referred to the related description of the signal processing module 6 described above.
  • the ultrasonic imaging method flow shown in FIG. 7 is provided in the ultrasonic gray scale imaging system shown in FIG. 1d.
  • step S110 is added.
  • the transmitting circuit 2 excites the probe to emit a plurality of sets of unfocused ultrasonic beams to the fluid, and also excites the probe to emit a focused ultrasonic beam to the scanning target containing the fluid.
  • the focused ultrasound beam here may also be a plurality of sets of focused ultrasound beams for obtaining ultrasound image data that is continuous in time.
  • step S210 is added, and the receiving circuit 4 and the beam combining module 5 receive the echoes of the focused ultrasonic beam returned from the scanning target to obtain a focused ultrasonic echo signal.
  • the receiving circuit 4 and the beam combining module 5 receive echoes of the plurality of sets of focused ultrasonic beams returning from the scanning target, and obtain a plurality of sets of focused ultrasonic echo signals for generating time-continuous ultrasonic image data.
  • step S320 in FIG. 7 After performing the above step S310 or after, step S320 in FIG. 7 is further included, after the signal detection and enhancement processing is performed on the focused ultrasound echo signal by the B mode signal processing module 11, the ultrasound image data is obtained, thereby Focusing on the ultrasound beam results in a sharper, better-imaged B-mode image.
  • the B-mode signal processing module 11 herein can refer to a conventional process of obtaining ultrasound image data using a focused ultrasound echo signal.
  • step S300 and step S320 the above step S400 is replaced with step S410, and the fluid display data is superimposed on the ultrasonic image data by the image processing module 7, and the data is converted to obtain the B mode ultrasonic image.
  • step S410 the fluid display data is superimposed on the ultrasonic image data by the image processing module 7, and the data is converted to obtain the B mode ultrasonic image.
  • display 8 displays the B-mode ultrasound image sequence described above in step 500 to present a visual effect of fluid dynamic flow.
  • the frame rate is relatively low when imaging with a focused ultrasound beam, but the ability of the focused ultrasound beam to be transmitted each time is concentrated, and imaging is only performed at the concentration of the power, thus obtaining the echo signal of the echo signal Higher than the ratio, it can be used to obtain better quality ultrasound images. Therefore, the present invention employs a combination of different ultrasonic beams in order to solve the problem that the gray-scale blood flow imaging has a distortion problem when displaying high-speed blood flow and the time rate of the ultrasonic image is too low. Focused ultrasound When the beam is beamed and unfocused, the focused ultrasound beam and the unfocused ultrasound beam can be alternately emitted. In order to keep the synchronization of the acquisition of the blood flow signal and the acquisition of the B-mode ultrasound image data as much as possible.
  • the transmitting circuit 2 can also excite the probe 1 to emit a plurality of focused ultrasonic beams of different emission angles to the scanning target containing the fluid, and to synthesize a pair of ultrasonic images based on the echoes of the focused ultrasonic beams of the plurality of different emission angles. Therefore, on the basis of ensuring the imaging frame rate, the obtained echo signal has a higher signal-to-noise ratio, and can be used to obtain better quality ultrasonic image data.
  • a plurality of unfocused ultrasound beams and focused ultrasound beams of different emission angles are alternately emitted.
  • the unfocused ultrasonic beam and the focused ultrasonic beam are alternately emitted to the scanning target containing the fluid in accordance with the above-described different emission angles.
  • multiple emission processes of unfocused ultrasound beams are transmitted to the fluid at different emission angles, and multiple emission processes of different focused emission angles are transmitted to the scanning target containing the fluid.
  • the cycle process is repeated a plurality of times to complete the unfocused ultrasonic beam and the focused ultrasound Multiple launches of the bundle.
  • the wall filter 10 can be added to the system architecture shown in FIG. 1d to form the ultrasonic grayscale image imaging system shown in FIG. 1c. Therefore, in step S410 in FIG. 7, the image processing module 7 superimposes the wall-filtered fluid display data on the ultrasound image data, and performs data conversion to obtain a B-mode ultrasound image sequence.
  • the image processing module 7 superimposes the wall-filtered fluid display data on the ultrasound image data, and performs data conversion to obtain a B-mode ultrasound image sequence.
  • the frequency of the beamformer output signal is typically moved by the demodulator to the baseband.
  • One way to accomplish this is to multiply the input signal by a complex sinusoidal signal.
  • f d is the amount of frequency shift that shifts the signal spectrum to the base band.
  • the demodulated signal is supplied to the signal processing module 6 and the image processing module 7, thereby converting the demodulated signal into corresponding image display data.
  • B mode gray level
  • this is a signal envelope that performs some additional processing, such as edge enhancement and logarithmic compression.
  • the image processing module needs to superimpose the data output from the signal processing module 7 and the B mode signal processing module 11 and convert it into a desired display image.
  • the image overlay module converts the acoustic image data from a polar coordinate (R-theta) sector format or a Cartesian coordinate linear array to a properly scaled Cartesian coordinate display pixel data at video rate.
  • the scanned converted acoustic data is then provided for display on display 8, which plots the time varying amplitude of the B-mode signal envelope in grayscale form. A corresponding scan line is displayed for each transmit beam.
  • fluid filtering is obtained after wall filtering, and the fluid signals are superimposed on the original image according to their positions, and the fluid flow effect can be displayed when played in time.
  • Different weights can be used in the superposition, or a suitable energy threshold can be used to first distinguish the location of the blood flow and the position without blood flow, and then display the filtered image only at the position of the blood flow.
  • the position without blood flow only shows the ultrasound image without filtering. Therefore, when the fluid display data is superimposed on the ultrasonic image data for display, the signal energy of the characterized fluid may be calculated for the wall-filtered fluid display data, and the ultrasonic image data obtained according to the focused ultrasonic beam is divided according to the energy threshold.
  • the fluid position region and the non-fluid position region superimpose the wall-filtered fluid display data on the fluid position region, and perform data conversion to obtain the B-mode ultrasound image sequence.
  • FIG. 6 or FIG. 7 is a schematic flow chart of an ultrasonic gray scale imaging method according to an embodiment of the present invention. It should be understood that although the various steps in the flowchart of FIG. 6 or FIG. 7 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. 6 or FIG.
  • the 7 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 execution order thereof is also It is not necessarily performed sequentially, but may be performed alternately or alternately with at least a portion of other steps or sub-steps or stages of other steps.
  • the above various embodiments are only described in the specific description for the implementation of the corresponding steps in FIG. 6 or FIG. 7. However, in the case where the logic is not contradictory, the above various embodiments can be combined with each other to form a new technical solution. The new technical solution is still within the scope of the present disclosure.
  • 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 ROM, disk, optical disk, server cloud space, including a number of instructions for causing a terminal device (which may be a mobile phone, a computer, a server, or a network device, etc.) to execute the system structure and method of various embodiments of the present invention.
  • a non-transitory computer readable storage carrier eg ROM, disk, optical disk, server cloud space, including a number of instructions for causing a terminal device (which may be a mobile phone, a computer, a server, or a network device, etc.) to execute the system structure and method of various embodiments of the present invention.
  • Various embodiments of the present invention are based on conventional gray-scale blood flow imaging techniques that employ unfocused beam emission, or a combination of unfocused and focused waves, with multiple scan lines or one shot at a time.
  • the display frame rate can be increased by 10 to 100 times compared with the conventional method.
  • Wall filtering of the collected signals can obtain blood flow signals.
  • each line needs to be scanned several times, and there is no filter transient problem.
  • Each of the scanned images obtained by the unfocused wave (for example, a plane beam) is continuously obtained, and the transient state is not considered in the filtering, and a plurality of images can be used for wall filtering to improve the signal-to-noise ratio, and the frame rate is such There will be no loss.

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Abstract

一种超声灰阶成像系统及方法。该系统包括：探头（1）；发射电路（2），用于激励探头（1）向流体发射非聚焦超声波束；接收电路（4）和波束合成模块（5），用于接收非聚焦超声波束的回波，获得非聚焦超声回波信号；信号处理模块（6），用于对多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；图像处理模块（7），用于将流体显示数据进行数据转换，获得B模式超声图像序列；显示器（8），用于显示B模式超声图像序列，以呈现流体动态流动的视觉效果。该系统采用非聚焦波或者平面波发射可以一次得到多条扫描线甚至整幅图像，可以极大的提高超声图像的时间分辨率，解决传统灰阶血流成像在显示高速血流时的失真问题。

Description

超声灰阶成像系统及方法 技术领域
本发明涉及超声成像系统，并且特别涉及采用灰阶图直接显示血流流动的一种超声血流成像方法。
背景技术
灰阶血流成像是将得到的普通B图的信号经过处理后采用灰阶图直接显示血流流动的一种超声血流成像方法。传统的灰阶血流成像采用的是聚焦波发射及逐线扫描方式，这样得到的血流精度和灵敏度较高，通过壁滤波增强血管内的流动信号，然后与组织信号相结合将血管壁和血管内红细胞的流动以灰阶图像同时显示出来。灰阶血流成像的动态图可以直观的显示流动血液与不流动组织的区别。然而，传统方法采用聚焦波发射以及逐线扫描，这会降低图像的时间分辨率，使灰阶血流成像的最大特点无法完全的发挥出来。当血流速度过快时，由于每条扫描线之间存在一个扫描时间间隔，因此会使灰阶血流成像的显示失真。
基于此现有技术中存在的问题，有待进一步地提高。
发明内容
基于此，有必要针对传统灰阶血流成像在显示高速血流时的失真问题，提供一种超声灰阶成像系统及方法。
本发明的一个实施例中，提供了一种超声灰阶成像系统，其包括：
探头；
发射电路，用于激励所述探头向流体发射多组非聚焦超声波束；
接收电路和波束合成模块，用于接收所述多组非聚焦超声波束的回波， 获得多组非聚焦超声回波信号；
信号处理模块，用于对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；
图像处理模块，用于将所述流体显示数据进行数据转换，获得B模式超声图像序列；及
显示器，用于显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
本发明的一个实施例中，提供了一种超声灰阶成像系统，其特征在于，所述系统包括：
探头；
发射电路，用于激励所述探头向包含流体的扫描目标发射多组非聚焦超声波束和聚焦超声波束；
接收电路和波束合成模块，用于接收所述多组非聚焦超声波束从所述流体上返回的回波，获得多组非聚焦超声回波信号，接收所述多组聚焦超声波束从所述扫描目标上返回的回波，获得多组聚焦超声回波信号；
信号处理模块，用于对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；
B模式信号处理模块，用于对所述多组聚焦超声回波信号进行信号检测、增强处理后，获得超声图像数据；
图像处理模块，用于将所述流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得B模式超声图像序列；及
显示器，用于显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
本发明的一个实施例中，提供了一种超声灰阶成像方法，所述方法包括：
向流体发射多组非聚焦超声波束；
接收所述多组非聚焦超声波束的回波，获得多组非聚焦超声回波信号；
对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体 显示数据；
将所述流体显示数据进行数据转换，获得B模式超声图像序列；
显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
本发明采用非聚焦波(如平面波或发散波)发射可以一次得到多条扫描线甚至整幅图像，这样可以极大的提高超声图像的时间分别率，解决传统灰阶血流成像在显示高速血流时的失真问题。
附图说明
图1a、图1b、图1c、图1d为本发明四种实施例的超声灰阶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的聚焦超声波束的示意图；
图5为本发明一个实施例中发散超声波束的示意图；
图6为本发明一个实施例的流程示意图；
图7为本发明另一个实施例的流程示意图；
图8为本发明的一个实施例中多角度发射平面超声波束的示意图；
图9为本发明的一个实施例中多角度发射非聚焦超声波束的示意图。
具体实施方式
图1a为本发明一个实施例的B模式超声灰阶成像系统的结构框图示意图。如图1a所示，该超声灰阶成像系统包括：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。上述信号处理模块6、图像处理模块7可以基于一个或多个处理器来实现。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，向包含流体的扫描目标发射超声波束，经一定延时后接收 从目标区域反射回来的带有流体信息的超声回波，并将此超声回波重新转换为电信号。扫描目标包括人体或者动物体内的器官、组织等中存在的血管或者生物体内其他其内有流体流动的脉管，图中未示出。流体包括人体或者动物体内血管、淋巴系统等组织中存在的液体、或者人体或者动物体内其他脉管内存在的流动液体，图中未示出。接收电路接收探头1转换生成的电信号，获得超声回波信号，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对超声回波信号进行聚焦延时、加权和通道求和等信号处理，然后将超声回波信号送入信号处理模块6进行信号检测、增强处理等相关的信号处理，获得包含流体信息的流体显示数据。图像处理模块7根据用户所需成像模式的不同，对流体显示数据进行不同的数据转换，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B模式超声图像，同时按照超声采集时间的顺序获得一系列的超声图像，本文称之为超声图像序列。图像处理模块7将生成的超声图像序列送入显示器8进行显示，就可以获得动态图像，从而展现流体动态流动的视觉效果。
探头1通常包括多个阵元的阵列。在每次发射超声波时，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，形成被发射到流体的合成超声波束，该合成超声波束的方向即为本文中所提到的发射方向。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。本文中提到的叠加，可以为两者的普通相加，还可以为按照一定的权重相加。
另外，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”。例如，图4为发射聚焦超声波束的示意图。这里，参与超声波的发射的阵元(图4中，仅仅探头1中的部分阵元参与了超声波的发射)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。
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，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。本文中，称这种无焦点的平面波为“平面超声波束”。
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”。如图5所示的发散超声波束。
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，合成的超声波束的发射方向与阵元排列平面的法线方向一致。例如，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，即为合成波束的偏转角度，改变上述时间延时，也就可以调整合成波束的偏转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偏转方向。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偏转角度。通过改变时延时间，可以调整角a的大小。
此外，还如图5所示，对于发散超声波束可以看做是虚焦点在探头后面的的非聚焦波，因此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可以调整虚焦点的位置，从而改变非聚焦波束的发射方向。同理，无论是平面超声波束、聚焦超声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“偏转角度”或者虚焦点位置，从而改变超声波束的发射角度，这里的合成波束可以为上文提到的平面超声波束、聚焦超声波束或发散超声波束等等，而本文中将平面超声波束和发散超声波束统称为非聚焦超声波束。
基于图1a提供的结构示意图，本发明采用非聚焦超声波束发射方案，一次发射可以得到多条扫描线或者得到一整幅图像。这样显示帧率与以前相比可以提高10-100倍。对收集的信号进行壁滤波可以得到血流信号，相比传统的通过聚焦超声波速的逐线扫描时每根线需要扫描若干次而需要克服滤波器的暂态问题，将更加有利于血流信号采集，这样可以极大的提高超声图像的时间分别率，解决传统灰阶血流成像在显示高速血流时的失真问题。在本文中将平面超声波束和发散超声波束统称为非聚焦超声波束。非聚焦超声波束，例如平面超声波束发射时每幅图像是连续得到的，滤波时不用考虑暂态，而且可以采用很多幅图做壁滤波从而提高信噪比，并且这样帧率还不会有损失。以下将结合各个附图详细说本发明的各种变形具体实施例。
基于图1a所示的系统架构示意图，图6给出了一种超声成像方法。
在步骤S 100中，发射电路2激励探头1向流体发射多组非聚焦超声波束。这里的非聚焦超声波束可以是平面超声波束或者发散超声波束。
多组非聚焦超声波束可以是按照时间先后顺序被激励获得的非聚焦超声波束。接收多组非聚焦超声波束从流体上返回的回波，获得多组超声回波信号，用以形成具有一定时间序列的图像序列，从而获得动态图像显示的源数据。
接收电路4和波束合成模块5接收一组非聚焦超声波束，可以获得一组非聚焦超声回波信号，而在本发明的其中一个实施例中，向流体发射的每组非聚焦超声波束中可以包括多个不同发射角度的非聚焦超声波束，接收电路4和波束合成模块5接收多个不同发射角度的非聚焦超声波束的回波，获得一组非聚焦超声回波信号包含的多次非聚焦超声回波信号；根据此多次非聚焦超声回波信号，对于同一空间位置处反射获得的非聚焦超声回波信号进行空间复合后，送入至信号处理模块6中。基于多个不同发射角度的非聚焦超声波束的回波，用于合成一副超声图像。从而能够在保证成像帧率的基础上，使获得的回波信号信噪比更高，可用以获得质量较好的超声图像数据。
此外，在本发明的其中一个实施例中，通过宽带宽的编码脉冲激励探头向流体发射非聚焦超声波束。使用宽带宽的编码脉冲，可以增强回波信号。这里提到的带宽是指该信号所包含的各种不同频率成分所占据的频率范围，而宽带宽是指频率范围为3兆至10兆甚至更宽。
在图1a的系统中，探头的每个阵元都使用基序列的编码序列作为驱动脉冲驱动(pulse)，序列中的每个脉冲通常被称为一片(CHIP)。基序列是采用N位的发射码进行相位编码的，以便产生N片的编码序列，这些序列被存储在发射序列存储器(未标出)中。从发射序列存储器中读出的每个编码序列在相应的发射激发期间控制发射电路2的激活。例如，在本发明的其中一个实施例中，发射电路2激励探头1向流体发射的每组非聚焦超声波束中可以包括多次向流体的发射，而每一次发射对应获得一次非聚焦超声回波信号，用以通过空间复合来获得更好显示效果的超声图像。然而这多次向流体发射的非聚焦超声波束可以是具有相同的发射角度，也可以具有不同的发射角度。当向流体发射的每组非聚焦超声波束中可以包括多个不同发射角度的非聚焦 超声波束、或包括分别沿多个不同发射角度的多次非聚焦超声波束时，按照发射角度的不同交替向流体发射非聚焦超声波束、或者按照发射角度的不同交替向流体发射多次非聚焦超声波束。图8所示，探头1分多个角度向流体发射平面超声波束，图中采用不同的线型来区分不同的发射角度。当然，在进行超声回波信号的空间复合时，图8中斜线区域的位置超声回波信号将会叠加。
当然，发射电路2激励探头1向流体发射的每组非聚焦超声波束中还可以包括多个不同位置的虚焦点下的非聚焦超声波束。按照虚焦点位置的不同交替向流体发射非聚焦超声波束、或按照虚焦点位置的不同交替向流体发射多次非聚焦超声波束。例如图9所示，虚焦点位置依次从A、B、到C的过程，对应的非聚焦超声波束分别用虚线、实线、点划线来表示。而在进行超声回波信号的空间复合时，图8中斜线区域的位置超声回波信号将会叠加。
在图1a所示的系统中，探头1通过供给的脉冲序列被激发N次，从而向流体发射N次波形相同的非聚焦超声波束。发射电路2驱动探头使得所产生的超声能量被引导或操纵在一行超声光束中覆盖整个扫描面，相比传统的灰阶血流成像系统中采用聚焦超波的方式，在帧率上更快、图像数据的获取速度更好，并且通过适当地调整发射焦点的时间延迟，使非聚焦超声波束可以被发射到所希望的偏转角度或虚焦点位置上。
在步骤S200中，接收电路4和波束合成模块5接收上述步骤S100中发射的多组非聚焦超声波束的回波，获得多组非聚焦超声回波信号。
每个超声能量短脉冲串(即超声波束)产生的超声回波信号都从沿着每个发射波束位于一连串射程的对象反射。超声回波信号分别由探头1检测，并且，超声回波信号在时间上某点处的幅度的采样表示在特定射程处发生的反射量。由于在反射点和探头1中每个阵元之间的传播通路的差异，超声回波信号不是被同时检测到，并且它们的振幅不相等。接收电路4接收探头1转换生成的电信号，获得相应的超声回波信号，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对超声回波信号进行聚焦延时、加权和通 道求和等处理。波束合成模块5跟踪发射的超声波束的方向，并对沿着每个超声波束的一连串射程处的超声回波信号进行抽样。波束合成模块5向每个超声回波信号赋予适当的时间延迟和接收变迹加权，并对这些信号求和从而得到合成的超声回波信号，该合成的超声回波信号准确地表示从沿着一次非聚焦超声波束的某个射程内的多个接收通道的对应信号之和。波束合成可以通过设置接收波束求和器来实现。对于同一周期内不同发射角度的超声波束的发射，其回波信号由接收电路获取。而接收电路中对应每个阵元的接收通道包括一个模一数转换器(未显示)。接收波束形成器存储器把适当的接收焦点时间延迟赋予每个接收的回波信号，并对这些回波信号求和而得到合成的回波信号，该合成信号准确地表示了从各个扫描位置反射的总超声能量。对于每个扫描位置，经过时间延迟的接收信号在接收波束求和器相加。
在本发明的其中一个实施例中，当发射电路2激励探头1向流体发射的每组非聚焦超声波束中包括多个不同发射角度的超声波束，则对应接收多个不同发射角度的超声波束的回波，获得多次非聚焦超声回波信号，根据此多次非聚焦超声回波信号，对于同一空间位置处反射获得的非聚焦超声回波信号进行空间复合可在波束合成之后。
在步骤S300中，信号处理模块6对上述非聚焦超声回波信号进行信号检测、增强处理后获得流体显示数据。这里的信号处理还可以包括对数压缩等等。而信号检测可以包括信号包络线检波处理。在此不详说明，可参见有关信号处理的常用方法。
在图1a所示的系统架构基础上，在本发明的其中一个实施例中，如图1b所示的超声成像系统还包括壁滤波器10。壁滤波器10对多组非聚焦超声回波信号进行壁滤波处理后获得滤波信号，而该滤波信号被送入到信号处理模块7进行信号检测、增强处理等相关信号处理后，以获得壁滤波后的流体显示数据。壁滤波器10可以采用传统的FIR、或者IIR、或者更复杂的比如线性回归滤波器，低秩滤波器等。而在本发明的其中一个实施例中，通过高通滤波器对非聚焦超声回波信号中每个相同位置在不同时刻的值进行滤波，即 将频率高的血流信号保留，滤掉频率低的非流体信号。通过壁滤波处理之后的回波信号会具有较高的信噪比，获得更好显示质量的超声图像数据。
在上述步骤S100中，向流体发射的每组非聚焦超声波束中包括多个不同发射角度的非聚焦超声波束，则接收电路4和波束合成模块5接收所述多个不同发射角度的非聚焦超声波束的回波，获得一组非聚焦超声回波信号中的多次非聚焦超声回波信号；根据所述多次非聚焦超声回波信号，对于同一空间位置处反射获得的非聚焦超声回波信号进行空间复合后，送入至壁滤波器10中。
此外，在本壁滤波器10包括两个部分，第一部分为提取基频分量，第二部分为使用高通滤波器抑制基波分量。多次非聚焦超声回波信号可以分别通过1个或多个壁滤波器进行壁滤波处理。
在步骤S400中，图像处理模块7将上述流体显示数据进行数据转换，获得B模式超声图像序列。
图像处理模块7对经过相关信号处理(包括边缘增强和对数压缩等)后的上述流体显示数据或者经过壁滤波后的流体显示数据进行数据转换，用以获得在显示器8中进行显示的超声图像序列。
经过相关信号处理之后的正交解调信号均包括I和Q两路，此I和Q两路可以分别进行壁滤波处理。对于滤波前和滤波后的I和Q两路解调数据在图1a和图1b所示的超声成像系统中，可以采用以下方式来获得相应的图像数据。
利用图像处理模块7对上述流体显示数据中的I和Q两路解调数据计算显示方差，根据显示方差获得区分非流体和流体图像区域，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，获得叠加有图像灰度和/或图像彩色信息的B模式超声图像序列，或者获得分区域叠加有图像灰度和/或图像彩色信息的B模式超声图像序列。方差大的表示流体，方差小的表示静止不动的非流体。而方差的计算公式可以如下所示：
显示方差Var表示为如下公式(1)：
Figure PCTCN2015088985-appb-000001
其中，i＝1，...，N，表示采样时刻。Ii表示第i时刻的I路解调数据，Qi表示第i时刻的Q路解调数据。
当然显示方差Var还可以表示为下述公式(2)。
Figure PCTCN2015088985-appb-000002
其中，K＝0，1，2，3，......。R(0)表示采用Leg-zero方法获得的流体信号能量，具体可以表示为下述公式(3)。
Figure PCTCN2015088985-appb-000003
依据上述公式(1)和公式(2)计算出的相应显示位置处的显示方差，按照显示方差的大小变化趋势映射不同的图像灰度和/或图像彩色信息，然而在图像处理模块7获得的B模式超声图像中叠加图像灰度和/或图像彩色信息，从而实现显示效果。例如，对整个像素图像中各个显示位置的显示方差做整体的图像灰度信息的映射，那么B模式超声图像中叠加图像灰度后获得的B模式超声图像序列在连续播放后显示的动态图像视觉效果时，全幅图呈现灰度效果，并在流体流动区域呈现动态的云朵状团簇体翻滚的视觉效果。还比如，对整个像素图像中各个显示位置的显示方差做整体的图像彩色信息的映射，那么B模式超声图像中叠加图像彩色信息后，获得的B模式超声图像序列在连续播放后显示的动态图像视觉效果时，全幅图呈现彩色对比效果，并在流体流动区域呈现动态的彩色云朵状团簇体翻滚的视觉效果。还可以是，按照显示方差阈值区分非流体和流体的图像位置，然后按照显示方差的大小变化趋势映射图像彩色信息，在流体的图像位置区域内叠加图像彩色信息。通过改变色相、明度、和饱和度中之一可以获得不同的图像色彩。按照显示 方差的大小变化趋势映射不同显示位置处的图像色彩，可以获得相应的图像彩色信息。
针对壁滤波处理之后的流体显示数据，还可以采用以下方式之一来进行数据转换或的B模式超声图像序列。
第一，对上述壁滤波后的流体显示数据中的I和Q两路解调数据计算显示方差，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，获得上述B模式超声图像序列。显示方差的计算方法参见上述公式(1)至公式(3)。
第二，对上述壁滤波后的流体显示数据中的I和Q两路解调数据进行极坐标系和直角坐标系的转换，并以图像灰度和/或图像彩色信息来映射信号包络的时变振幅趋势，获得上述B模式超声图像序列。例如，直接显示血流图像Flow_image(x，z)，B模式超声图像序列的转换方式参见下述公式：
Figure PCTCN2015088985-appb-000004
直接得到图像上每个点的值，方法一类似常规的B图成像方法。
第三，根据上述壁滤波后的流体显示数据中的I和Q两路解调数据，计算表征流体的信号能量，以图像灰度和/或图像彩色信息来映射信号能量的大小，获得上述B模式超声图像序列。例如，采用Leg-zero方法，显示血流信号能量，计算方式可参见上述公式(3)。
当图像数据只有一路数据时，可以先对图像的RF数据沿深度(Depth)做Hilbert变换，得到I和Q两路数据，然后再做壁滤波处理，从用上述三种方法之一得到B模式超声图像序列。
上述图像处理模块还用于根据上述壁滤波后的流体显示数据，计算表征流体的信号能量，基于能量阈值划分上述超声图像数据中流体位置区域和非流体位置区域，将上述流体位置区域上叠加上述壁滤波后的流体显示数据，进行数据转换后获得上述B模式超声图像序列。这样可以使获得的图像数据能够获得更加清晰、和准确的流体图像信息。
依据上述公式(3)计算出的相应显示位置处的信号能量，按照信号能量的大小变化趋势映射不同的图像灰度和/或图像彩色信息，然而在图像处理模块7获得的B模式超声图像中叠加图像灰度和/或图像彩色信息，从而实现显示效果。例如，对整个像素图像中各个显示位置的信号能量做整体的图像灰度信息的映射，那么B模式超声图像中叠加图像灰度后获得的B模式超声图像序列在连续播放后显示的动态图像视觉效果时，全幅图呈现灰度效果，并在流体流动区域呈现动态的云朵状团簇体翻滚的视觉效果。还比如，对整个像素图像中各个显示位置的信号能量做整体的图像彩色信息的映射，那么B模式超声图像中叠加图像彩色信息后，获得的B模式超声图像序列在连续播放后显示的动态图像视觉效果时，全幅图呈现彩色对比效果，并在流体流动区域呈现动态的彩色云朵状团簇体翻滚的视觉效果。还可以是，按照信号能量阈值区分非流体和流体的图像位置，然后按照信号能量的大小变化趋势映射图像彩色信息，在流体的图像位置区域内叠加图像彩色信息。通过改变色相、明度、和饱和度中之一可以获得不同的图像色彩。按照信号能量的大小变化趋势映射不同显示位置处的图像色彩，可以获得相应的图像彩色信息。
在步骤S500中，显示器8显示B模式超声图像序列，以呈现流体动态流动的视觉效果。例如，在显示器8的界面上获得在流体流动区域呈现动态的灰色或彩色云朵状团簇体翻滚的视觉效果。
在上述各个实施例的基础上，如图1c所示的超声成像系统中还包括B模式信号处理模块11，上述多组非聚焦超声回波信号分两路输出，一路依次输入至上述壁滤波器10、信号处理模块6后获得壁滤波后的流体显示数据，另一路输入至上述B模式信号处理模块11进行信号检测、增强处理后，获得超声图像数据；上述图像处理模块7还用于将上述壁滤波后的流体显示数据叠加在上述超声图像数据上，并进行数据转换后获得上述B模式超声图像序列。B模式信号处理模块11的信号过程可参见上述有关信号处理模块6的相关说明。
另外，在本发明的另一个实施例中，为了通过聚焦超声波束获得更加高 质量的背景图像。因此，基于图1d所示的超声灰阶成像系统中提供了图7所示的超声成像方法流程。
基于上述步骤S100，增加步骤S110，发射电路2除了激励探头向流体发射多组非聚焦超声波束，还激励上述探头向包含流体的扫描目标发射聚焦超声波束。而这里的聚焦超声波束也可以是多组聚焦超声波束，用以获得按照时间连续的超声图像数据。
然后，基于上述步骤S200，增加步骤S210，接收电路4和波束合成模块5接收聚焦超声波束从扫描目标上返回的回波，获得聚焦超声回波信号。例如，接收电路4和波束合成模块5接收多组聚焦超声波束从扫描目标上返回的回波，获得多组聚焦超声回波信号，用以生成时间连续的超声图像数据。
于是，在执行上述步骤S310的同时或者之后，还包括图7中的步骤S320，利用B模式信号处理模块11对上述聚焦超声回波信号进行信号检测、增强处理后，获得超声图像数据，从而根据聚焦超声波束可获得更加清晰、成像效果更好的B模式图像。这里的B模式信号处理模块11可参照常规的利用聚焦超声回波信号获得超声图像数据的过程。
最后，在图7中基于上述步骤S300和步骤S320，上述步骤S400替换为步骤S410，利用图像处理模块7将上述流体显示数据叠加在上述超声图像数据上，并进行数据转换后获得B模式超声图像序列；在步骤500中显示器8显示上述B模式超声图像序列，以呈现流体动态流动的视觉效果。
上述过程中与图6中各个方法步骤相似或相同的部分可参见前文说明，而有关聚焦波束的回波信号的处理过程参见常规的有关聚焦超声波束的处理过程。
在本发明的其中一个实施例中，使用聚焦超声波束成像时帧率相对较低，但是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，可用以获得质量较好的超声图像。因此，本发明采用了不同超声波束的组合方式，以求解决灰阶血流成像在显示高速血流时的失真问题和超声图像的时间分别率过低的问题。上述同时存在聚焦超声波 束和非聚焦超声波束时，可以令聚焦超声波束与非聚焦超声波束交替发射。以求在血流信号的获取和B模式超声图像数据的获取上时间保持尽可能同步。
此外，发射电路2还可以激励探头1向包含流体的扫描目标发射多个不同发射角度的聚焦超声波束，基于多个不同发射角度的聚焦超声波束的回波，用于合成一副超声图像。从而能够在保证成像帧率的基础上，使获得的回波信号信噪比更高，可用以获得质量较好的超声图像数据。
在本发明的另一个实施例中，多个不同发射角度的非聚焦超声波束和聚焦超声波束交替发射。例如，按照上述发射角度的不同交替向包含流体的扫描目标发射非聚焦超声波束和聚焦超声波束。在同一个发射周期内，先执行不同发射角度向流体发射非聚焦超声波束的多次发射过程，再执行不同发射角度向包含流体的扫描目标发射聚焦超声波束的多次发射过程。或者，沿同一发射角度依次执行向流体发射非聚焦超声波束的发射过程和向包含流体的扫描目标发射聚焦超声波束的发射过程后，多次重复这一周期过程从而完成非聚焦超声波束和聚焦超声波束的多次发射。
当然，在图1d所示的系统架构上还可以增加壁滤波器10形成图1c所示的超声灰阶图像成像系统。因此，在图7中的步骤S410中，利用图像处理模块7将经过壁滤波处理后的流体显示数据叠加在上述超声图像数据上，并进行数据转换后获得B模式超声图像序列。有关利用壁滤波器10对非聚焦超声波束的回波信号进行处理的过程参见前文有关步骤S300的说明，在此不再累述。
在上述图1c和图1d所示的各个实施例的超声成像系统中，波束合成器输出信号的频率通常由解调器移动到基频带。完成这一过程的一种途径是将输入信号乘以复数正弦信号
Figure PCTCN2015088985-appb-000005
这里fd是将信号频谱移到基频带的频率移动量。解调信号被提供给信号处理模块6和图像处理模块7，从而将解调信号转换成相应的图像显示数据。在B模式(灰度级)中，这是进行了一些额外处理，比如边缘增强和对数压缩的信号包络。图像处理模块需要将来自信号 处理模块7和B模式信号处理模块11输出的数据进行叠加后，转换成所希望的显示图像。特别是，图像叠加模块将声学图象数据从极坐标(R-θ)扇形格式或者笛卡尔坐标线性阵列转换为在视频速率下的恰当定标的笛卡尔坐标显示像素数据。然后经过扫描转换的声学数据被提供在显示器8上进行显示，该显示器将B-mode信号包络的时变振幅以灰度级的形式绘制出来。对于每个发射波束显示一个相应的扫描线。
因此，在本发明的其中一个实施例中，实现壁滤波后得到的是流体信号，将这些流体信号按照其位置叠加在原图像上，按时间播放时就可以显示出流体的流动效果。而在叠加时可以采用不同的权重，也可以基于血流能量采用一个合适的能量阈值首先分辨出有血流的位置和无血流的位置，然后将有血流的位置只显示滤波后的图像，无血流的位置只显示没有滤波的超声图像。所以，上述流体显示数据叠加在上述超声图像数据上进行显示时，可以针对上述壁滤波后的流体显示数据，计算表征流体的信号能量，基于能量阈值划分上述根据聚焦超声波束获得的超声图像数据中流体位置区域和非流体位置区域，将上述流体位置区域上叠加上述壁滤波后的流体显示数据，进行数据转换后获得上述B模式超声图像序列。
图6或图7为本发明实施例的超声灰阶成像方法的流程示意图。应该理解的是，虽然图6或图7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6或图7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轮流或者交替地执行。以上各个实施例在具体说明中仅只针对图6或图7中相应步骤的实现方式进行了阐述，然而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的系统结构和方法。
本发明的各个实施例中基于传统的灰阶血流成像技术进行了改进，其采用非聚焦波发射，或者非聚焦波与聚焦波的组合发射方案，一次发射可以得到多条扫描线或者得到一整幅图像，通过这样的采集方式可以提高显示帧率，相比传统方式可以提高10-100倍。对收集的信号进行壁滤波可以得到血流信号，相比传统的逐线扫描每根线需要扫描若干次的扫描方式，不存在滤波器暂态问题。通过非聚焦波(例如平面波束)获得的扫描图像的每一幅都是连续得到的，滤波时不用考虑暂态，而且还可以采用很多幅图做壁滤波从而提高信噪比，并且这样帧率还不会有损失。
以上实施例仅表达了几种实施方式，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 (29)

  1. 一种超声灰阶成像系统，其特征在于，所述系统包括：
    探头；
    发射电路，用于激励所述探头向流体发射多组非聚焦超声波束；
    接收电路和波束合成模块，用于接收所述多组非聚焦超声波束从所述流体上返回的回波，获得多组非聚焦超声回波信号；
    信号处理模块，用于对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；
    图像处理模块，用于将所述流体显示数据进行数据转换，获得B模式超声图像序列；及
    显示器，用于显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
  2. 根据权利要求1所述的系统，其特征在于，所述系统还包括：
    壁滤波器，用于对所述多组非聚焦超声回波信号进行壁滤波处理后，送入到所述信号处理模块，以获得壁滤波后的流体显示数据。
  3. 根据权利要求2所述的系统，其特征在于，所述系统还包括：B模式信号处理模块，
    所述多组非聚焦超声回波信号分两路输出，一路依次输入至所述壁滤波器、信号处理模块后获得壁滤波后的流体显示数据，另一路输入至所述B模式信号处理模块进行信号检测、增强处理后，获得超声图像数据；
    所述图像处理模块还用于将所述壁滤波后的流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列。
  4. 根据权利要求1或2所述的系统，其特征在于，所述系统中，向流体发射的每组非聚焦超声波束中包括多个不同发射角度的非聚焦超声波束，所述接收电路和波束合成模块接收所述多个不同发射角度的非聚焦超声波束的回波，获得一组非聚焦超声回波信号中的多次非聚焦超声回波信号；根据所 述多次非聚焦超声回波信号，对于同一空间位置处反射获得的非聚焦超声回波信号进行空间复合后，送入至所述信号处理模块或壁滤波器中。
  5. 根据权利要求1所述的系统，其特征在于，所述图像处理模块还用于对所述流体显示数据中的I和Q两路解调数据计算显示方差，根据所述显示方差阈值区分非流体和流体图像区域，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，获得叠加有图像灰度和/或图像彩色信息的B模式超声图像序列、或者获得分区域叠加有图像灰度和/或图像彩色信息的B模式超声图像序列。
  6. 根据权利要求2或3所述的系统，其特征在于，所述图像处理模块还用于通过以下方式之一进行数据转换后获得所述B模式超声图像序列：
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据计算显示方差，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，获得所述B模式超声图像序列；
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据进行极坐标系和直角坐标系的转换，并以图像灰度和/或图像彩色信息来映射信号包络的时变振幅趋势，获得所述B模式超声图像序列；或者，
    根据所述壁滤波后的流体显示数据中的I和Q两路解调数据，计算表征流体的信号能量，以图像灰度和/或图像彩色信息来映射信号能量的大小，获得所述B模式超声图像序列。
  7. 根据权利要求1所述的系统，其特征在于，通过宽带宽的编码脉冲激励所述探头向流体发射非聚焦超声波束。
  8. 根据权利要求3所述的系统，其特征在于，所述图像处理模块还用于根据所述壁滤波后的流体显示数据，计算表征流体的信号能量，基于能量阈值划分所述超声图像数据中流体位置区域和非流体位置区域，将所述流体位置区域上叠加所述壁滤波后的流体显示数据，进行数据转换后获得所述B模式超声图像序列。
  9. 根据权利要求1所述的系统，其特征在于，向流体发射的每组非聚焦 超声波束中包括多个不同位置的虚焦点下的非聚焦超声波束。
  10. 一种超声灰阶成像系统，其特征在于，所述系统包括：
    探头；
    发射电路，用于激励所述探头向包含流体的扫描目标发射多组非聚焦超声波束和聚焦超声波束；
    接收电路和波束合成模块，用于接收所述多组非聚焦超声波束从所述流体上返回的回波，获得多组非聚焦超声回波信号，接收所述多组聚焦超声波束从所述扫描目标上返回的回波，获得多组聚焦超声回波信号；
    信号处理模块，用于对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；
    B模式信号处理模块，用于对所述多组聚焦超声回波信号进行信号检测、增强处理后，获得超声图像数据；
    图像处理模块，用于将所述流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得B模式超声图像序列；及
    显示器，用于显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
  11. 根据权利要求10所述的系统，其特征在于，所述系统还包括：
    壁滤波器，用于对所述多组非聚焦超声回波信号进行壁滤波处理后，送入到所述信号处理模块，以获得壁滤波后的流体显示数据；
    所述图像处理模块用于将所述壁滤波后的流体显示数据叠加
    在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列。
  12. 根据权利要求10或11所述的系统，其特征在于，所述系统中，向包含流体的扫描目标发射的每组非聚焦超声波束中包括多个不同发射角度的非聚焦超声波束，所述接收电路和波束合成模块接收所述多个不同发射角度的非聚焦超声波束的回波，获得一组非聚焦超声回波信号中的多次非聚焦超声回波信号；根据所述多次非聚焦超声回波信号，对于同一空间位置处反射 获得的非聚焦超声回波信号进行空间复合后，送入至所述信号处理模块或壁滤波器中。
  13. 根据权利要求10所述的系统，其特征在于，所述图像处理模块还用于对所述流体显示数据中的I和Q两路解调数据计算显示方差，根据所述显示方差区分非流体和流体图像区域，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，用以获得叠加有图像灰度和/或图像彩色信息的B模式超声图像序列、或者分区域叠加有图像灰度和/或图像彩色信息的B模式超声图像序列。
  14. 根据权利要求10或11所述的系统，其特征在于，所述图像处理模块还用于通过以下方式之一进行数据转换后获得所述B模式超声图像序列：
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据计算显示方差，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，用以获得所述B模式超声图像序列；
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据进行极坐标系和直角坐标系的转换，并以图像灰度和/或图像彩色信息来映射信号包络的时变振幅趋势，用以获得所述B模式超声图像序列；或者，
    根据所述壁滤波后的流体显示数据中的I和Q两路解调数据，计算表征流体的信号能量，以图像灰度和/或图像彩色信息来映射信号能量的大小，用以获得所述B模式超声图像序列。
  15. 根据权利要求10所述的系统，其特征在于，通过宽带宽的编码脉冲激励所述探头发射非聚焦超声波束和聚焦超声波束。
  16. 根据权利要求10所述的系统，其特征在于，所述非聚焦超声波束和聚焦超声波束交替发射。
  17. 根据权利要求11所述的系统，其特征在于，所述图像处理模块还用于根据所述壁滤波后的流体显示数据，计算表征流体的信号能量，基于能量阈值划分所述超声图像数据中流体位置区域和非流体位置区域，将所述流体位置区域上叠加所述壁滤波后的流体显示数据，进行数据转换后获得所述B 模式超声图像序列。
  18. 根据权利要求10所述的系统，其特征在于，向流体发射的每组非聚焦超声波束中包括多个不同位置的虚焦点下的非聚焦超声波束。
  19. 一种超声灰阶成像方法，所述方法包括：
    向流体发射多组非聚焦超声波束；
    接收所述多组非聚焦超声波束的回波，获得多组非聚焦超声回波信号；
    对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据；
    将所述流体显示数据进行数据转换，获得B模式超声图像序列；
    显示所述B模式超声图像序列，以呈现流体动态流动的视觉效果。
  20. 根据权利要求19所述的方法，其特征在于，所述对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据的步骤包括：
    对所述多组非聚焦超声回波信号进行壁滤波处理，获得滤波信号；
    将所述滤波信号进行信号检测、增强处理后，获得壁滤波后的流体显示数据。
  21. 根据权利要求19所述的方法，其特征在于，所述对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得流体显示数据的步骤包括：
    对所述多组非聚焦超声回波信号经过壁滤波处理、信号检测、增强处理后，获得壁滤波后的流体显示数据；
    对所述多组非聚焦超声回波信号进行信号检测、增强处理后，获得超声图像数据；
    所述将所述流体显示数据进行数据转换，获得B模式超声图像序列的步骤包括：
    将所述壁滤波后的流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列。
  22. 根据权利要求19所示的方法，其特征在于，向流体发射的每组非聚焦超声波束中包括多个不同发射角度的非聚焦超声波束，
    所述接收所述多组非聚焦超声波束的回波，获得多组非聚焦超声回波信号的步骤中包括：接收所述多个不同发射角度的非聚焦超声波束的回波，获得一组非聚焦超声回波信号中的多次非聚焦超声回波信号；
    根据所述多次非聚焦超声回波信号，对于同一空间位置处反射获得的非聚焦超声回波信号进行空间复合后，进行信号检测、增强处理或进行壁滤波处理。
  23. 根据权利要求20所示的方法，其特征在于，所述方法还包括：
    向包含流体的扫描目标发射聚焦超声波束；
    接收所述聚焦超声波束从所述扫描目标上返回的回波，获得聚焦超声回波信号；
    对所述聚焦超声回波信号进行信号检测、增强处理后，获得超声图像数据；
    所述将所述流体显示数据进行数据转换，获得B模式超声图像序列的步骤包括：
    将所述壁滤波后的流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列。
  24. 根据权利要求19所述的方法，其特征在于，所述将所述流体显示数据进行数据转换，获得B模式超声图像序列的步骤包括：
    对所述流体显示数据中的I和Q两路解调数据计算显示方差，根据所述显示方差获得区分非流体和流体图像区域，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，用以获得叠加有图像灰度和/或图像彩色信息的B模式超声图像序列、或者获得分区域叠加有图像灰度和/或图像彩色信息的B模式超声图像序列。
  25. 根据权利要求21或23所述的方法，其特征在于，所述将所述壁滤波后的流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列的步骤包括：
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据计算显示方 差，以图像灰度和/或图像彩色信息映射显示方差的变化趋势，获得所述B模式超声图像序列；
    对所述壁滤波后的流体显示数据中的I和Q两路解调数据进行极坐标系和直角坐标系的转换，并以图像灰度和/或图像彩色信息来映射信号包络的时变振幅趋势，获得所述B模式超声图像序列；或者，
    根据所述壁滤波后的流体显示数据中的I和Q两路解调数据，计算表征流体的信号能量，以图像灰度和/或图像彩色信息来映射信号能量的大小，获得所述B模式超声图像序列。
  26. 根据权利要求19或23所述的方法，其特征在于，通过宽带宽的编码脉冲激励所述探头发射非聚焦超声波束和/或聚焦超声波束。
  27. 根据权利要求23所述的方法，其特征在于，所述非聚焦超声波束和聚焦超声波束交替发射。
  28. 根据权利要求23所述的方法，其特征在于，所述将所述壁滤波后的流体显示数据叠加在所述超声图像数据上，并进行数据转换后获得所述B模式超声图像序列的步骤还包括：
    根据所述壁滤波后的流体显示数据，计算表征流体的信号能量，基于能量阈值划分所述超声图像数据中流体位置区域和非流体位置区域，将所述流体位置区域上叠加所述壁滤波后的流体显示数据，进行数据转换后获得所述B模式超声图像序列。
  29. 根据权利要求19所述的方法，其特征在于，向流体发射的每组非聚焦超声波束中包括多个不同位置的虚焦点下的非聚焦超声波束。
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