JP4713862B2 - Ultrasonic diagnostic equipment - Google Patents

Description

The present invention relates to ultrasonic diagnostic equipment which utilizes the ultrasonic diagnosing such blood flow or tissue motion.

  Usually, when quantifying blood flow or tissue motion, it is necessary to set a region of interest (ROI) according to the direction of motion.

  For example, in blood flow measurement, it has been proposed to obtain a blood flow using a flow velocity profile with an angle corrected and an ROI in a direction orthogonal to the blood vessel. In PW Doppler, Doppler angle correction is performed by designating the direction of blood flow. Alternatively, as an automatic angle correction method, a method of determining a blood flow direction from profile information has been proposed.

  It has also been proposed to move (track) the position of the region of interest in accordance with the movement of the tissue (see, for example, Patent Document 1). Patent Document 1 proposes to automatically move the ROI position in accordance with the amount of movement of the tissue for each image frame input in time series in order to prevent the ROI position from being shifted due to the movement of the tissue. .

As a method for detecting the motion speed of a living body, there is a method using the Doppler method, which detects a motion component in the beam direction. A method for detecting a two-dimensional or three-dimensional vector motion has also been proposed.
Japanese Patent Application No. 8-310820

  In the above-described conventional technique, for example, in blood flow measurement, angle correction is performed according to the direction of blood flow on the image, or ROI is set in an appropriate direction with respect to the direction of blood flow on the image. There is a need to. Conventionally, information input for the angle correction as described above and setting of the ROI depend on manual operation. For this reason, there is a problem that the operator's work is complicated, the error is large, and the reproducibility cannot be ensured.

  In addition, in order to automatically acquire information for angle correction as described above, a technique for calculating a blood flow direction by image processing has also been proposed. However, there is a problem in that very advanced calculation processing is required. there were.

The present invention has been made in consideration of such circumstances, and the object of the present invention is to provide a medical image analysis apparatus capable of automatically setting a region of interest in an appropriate direction with respect to the motion direction of a moving body. Is to provide.

Transmission / reception means for transmitting / receiving a Doppler beam at two different angles with respect to a plurality of measurement sites within a measurement range, and two received at each of the two angles for one measurement site of the plurality of measurement sites Calculation means for executing a process for calculating a velocity vector of a moving body at the one measurement site based on two velocity vectors obtained individually from a Doppler beam, and the calculation for the plurality of measurement sites. Generating means for generating an image representing the motion state of the moving body within the measurement range based on a plurality of velocity vectors respectively calculated by the means, and setting a rectangular region of interest in the image according to a user operation A setting means for performing an operation on the velocity vector of the moving body within the region of interest set by the setting means. And changing means for one side of the band to change the orientation of the region of interest so as to be orthogonal, and means for calculating a feature quantity of the moving body with respect to the orientation has changed the ROI to the ultrasonic diagnostic apparatus.

According to the present invention, it is possible to provide a medical image analysis apparatus capable of automatically setting a region of interest in an appropriate direction with respect to a moving direction of a moving body.

  Hereinafter, an embodiment of the present invention will be described with reference to the drawings.

FIG. 1 is a diagram showing a configuration of an ultrasonic diagnostic apparatus according to this embodiment.
As shown in FIG. 1, the ultrasonic diagnostic apparatus includes an ultrasonic probe 1, a transmission system circuit 2, a reception system circuit 3, a B mode processing circuit 4, a CFM mode processing circuit 5, a TDI mode processing circuit 6, and a PWD mode processing circuit 7. , Digital scan converter (DSC) 8, speed vector calculator 9, image memory unit 10, average vector calculator 11, ROI generator 12, feature quantity calculator 13, data synthesizer 14, frame synthesizer 15, display unit 16 And a CPU 17.

  The ultrasonic probe 1 includes an array-type piezoelectric vibrator disposed at the tip thereof. An array-type vibrator has a plurality of piezoelectric elements arranged in a straight line and the arrangement direction is set as a scanning direction, and each of the plurality of piezoelectric elements forms a transmission / reception channel. The ultrasonic probe 1 can perform bidirectional signal conversion between the ultrasonic signal and the electric signal. The ultrasonic probe 1 is connected to a transmission system circuit 2 and a reception system circuit 3. The ultrasonic probe 1 is configured as an independent unit separated from the main body of the ultrasonic diagnostic apparatus, and may be connected to the transmission system circuit 2 and the reception system circuit 3 as necessary by a connector or the like. . In this case, the main body of the ultrasonic diagnostic apparatus and the ultrasonic probe 1 may be distributed as a set or may be distributed individually.

  Although not shown, the transmission system circuit 2 includes a pulse generator and a transmission circuit. The pulse generator generates a reference rate pulse. The transmission circuit delays the reference rate pulse for each channel to generate a drive pulse. The transmission system circuit 2 supplies the drive pulse for each channel to the plurality of transducers of the ultrasonic probe 1. The transmission delay time of the drive pulse is controlled for each channel and is repeatedly supplied for each rate frequency. In response to the supply of the drive pulse, an ultrasonic pulse is emitted from each transducer. The ultrasonic pulse propagates through the subject P, forms a transmission beam according to the controlled transmission delay time, and reflects a part of the ultrasonic pulse at an interface having different acoustic impedances to become an echo signal. Part or all of the returned echo signal is received by one or more transducers and converted into a corresponding electrical signal.

  Although not shown, the reception system circuit 3 includes a preamplifier, a delay circuit, and an adder. A preamplifier is provided for each channel connected to each transducer of the ultrasonic probe 1. The preamplifier amplifies the analog signal of the electrical quantity corresponding to the echo signal received by the ultrasonic probe 1 for each channel. The delay circuit is connected to each preamplifier. The delay circuit delay-controls an analog signal output from the preamplifier for reception focus for each channel. The adder adds the analog signals for each channel after delay control. As a result, a reception beam having a focus point determined according to the control of the reception delay time is formed in calculation, and desired directivity is obtained.

  The output terminal of the reception system circuit 3 is connected to the B mode processing circuit 4, the CFM mode processing circuit 5, the TDI mode processing circuit 6 and the PWD mode processing circuit 7, respectively.

  The B-mode processing circuit 4 is responsible for creating B-mode black and white tomographic image data, and includes a logarithmic amplifier, an envelope detector, and an A / D converter (not shown). The logarithmic amplifier logarithmically compresses and amplifies the echo signal phased and added by the reception system circuit 3. The envelope detector detects the envelope of the output signal of the logarithmic amplifier. The A / D converter converts the output signal of the envelope detector into a digital signal. The output of the A / D converter is output from the B mode processing circuit 4 as B mode image data.

  The CFM mode processing circuit 5 is configured by a conventionally known circuit that two-dimensionally detects blood flow information by a color flow mapping (CFM: one type of color Doppler tomography) mode. Although not specifically shown, the CFM mode processing circuit 5 includes an orthogonal phase detector, an A / D converter, an MTI filter, and an autocorrelator, and an average that performs an operation based on the autocorrelation output. A speed calculator, a distributed calculator, and a power calculator are provided. The quadrature detector detects a Doppler signal from the received echo signal. The A / D converter converts the detected Doppler signal into digital data. The MTI filter temporarily stores digital data output from the A / D converter in a built-in frame memory.

  In the CFM mode, in order to obtain blood flow information, the same cross section is scanned a plurality of times. Therefore, the above frame memory stores Doppler data having three dimensions: a beam scan direction, an ultrasonic beam direction, and a scan frequency direction. The The MTI filter includes a high-pass filter on the reading side of the frame memory. Therefore, the Doppler component of the blood flow echo is removed by removing the Doppler component of the tissue echo from each of the plurality of Doppler data sequences in the scan frequency direction corresponding to each pixel position in the image data having three dimensions. Extracted into The self-computation unit analyzes the average Doppler frequency of the high-pass filtered Doppler data sequence. Based on the above average Doppler frequency, the average velocity calculator, dispersion calculator, and power calculator respectively calculate the average blood flow velocity at each sample point of the scan section, the variance value of the blood flow velocity distribution, and the echo signal from the blood flow. Each power value is calculated. These pieces of calculation information are output from the CFM mode processing circuit 5 as color Doppler information.

  The TDI mode processing circuit 6 detects tissue motion information two-dimensionally by tissue Doppler imaging (TDI: a type of color Doppler tomography). The configuration of the TDI mode processing circuit 6 is roughly the same as that of the CFM mode processing circuit 5 described above, but the characteristics of the filter circuit provided in the MTI filter can be extracted from the Doppler component of the echo signal from the tissue such as the myocardium. It is set as follows. That is, the characteristic is that there is a difference in intensity and Doppler shift frequency (motion speed) between the tissue echo signal and the blood flow echo signal. The intensity of the tissue echo signal is greater than that of the blood flow, but the Doppler shift frequency (ie velocity) is usually small. For this reason, the filter circuit mounted on the MTI filter is configured as a low-pass filter that can extract the low-frequency Doppler shift frequency. Other configurations are the same as those of the CFM mode processing circuit 5.

  The PWD mode processing circuit 7 has a function of generating Doppler spectrum data based on a pulse Doppler (PWD) method. Specifically, a quadrature phase detector, a sample hold circuit, a band filter, an A / D converter, an FFT, and the like are provided.

  The DSC 8 reconstructs an image from signals output from the processing circuits 4 to 7 and given through the velocity vector calculator 9. The DSC 8 also has a function of generating an image for simultaneously displaying information acquired in a plurality of modes by using a built-in memory (not shown).

  The velocity vector calculator 9 calculates a two-dimensional vector based on output signals from the CFM mode processing circuit 5 and the TDI mode processing circuit 6. The two-dimensional vector calculated by the velocity vector calculator 9 is stored in the image memory unit 10 and input to the average vector calculator 11 via the image memory unit 10. The average vector calculator 11 calculates a spatial average flow velocity vector based on the two-dimensional vector calculated by the velocity vector calculator 9 or the two-dimensional vector stored in the image memory unit 10. Further, the average vector calculator 11 has a function of generating graphic data indicating the direction and speed of blood flow in color based on the calculated average flow velocity vector. The average vector calculator 11 has a function of outputting information for determining the ROI direction based on the calculated average flow velocity vector.

  The ROI generator 12 sets ROI and angle markers. The ROI generator 12 has a function of automatically setting the inclination of the ROI and the angle marker based on the information for determining the ROI direction output from the average vector calculator 11.

  The feature amount calculator 13 calculates a feature amount in the ROI set by the ROI generator 12.

  The data synthesizer 14 is a frame capable of displaying the blood flow graphic data output from the average vector calculator 11 and the ROI and angle marker graphic data output from the ROI generator 12 at each designated position on the screen. Convert to data.

  The frame synthesizer 15 synthesizes pixels of the frame of the image data output from the DSC 8 and the frame data output from the data synthesizer 14 to generate a display image in which the reconstructed image is combined with the graphic. The display unit 16 displays the display image generated by the frame synthesizer 15.

  The CPU 17 performs control for realizing a change in pulse transmission sequence and a change in mode.

  Next, the operation of the ultrasonic diagnostic apparatus configured as described above will be described. Note that the operation of acquiring information about the subject using ultrasound and the operation of reconstructing an image from the acquired information are the same as those of conventional ultrasonic diagnostic apparatuses, and thus the description thereof is omitted here. . Since the feature of the present embodiment lies in the calculation of the blood flow velocity vector, that is, the flow velocity vector and the setting of the ROI slope, these processes will be described in detail below.

  In order to calculate the flow velocity vector, the ultrasonic diagnostic apparatus of this embodiment performs inverted transmission / reception. That is, the ultrasonic diagnostic apparatus of this embodiment performs Doppler beam scanning in two beam directions as shown in FIG. The beam direction may be reversed every frame or may be reversed every rate. The Doppler beam scan in such a form can be used for the B mode / color independent scan disclosed in Japanese Patent Application No. 62-69001.

  As shown in FIG. 2, a plurality of Doppler beams in each Doppler beam scan are parallel to each other. The scan planes in different Doppler beam scans are the same plane. The crossing angle between Doppler beams in different Doppler beam scans is always substantially constant.

  In such a region where two Doppler beam scans overlap, Doppler signals from different directions are obtained, and a blood flow velocity vector can be calculated based on these signals. In the present embodiment, a region indicated by hatching in FIG. 3 is a blood flow velocity vector measurement region.

  The velocity vector calculator 9 calculates a blood flow velocity vector for each measurement site in the measurement region based on Doppler signals respectively obtained by two Doppler beam scans for this measurement site.

FIG. 4 is a diagram showing the principle of calculating the blood flow velocity vector.
Assume that velocity vectors V1 and V2 are obtained from two Doppler signals relating to the measurement site in FIG. The position of the tip of the blood velocity vector is obtained as the intersection of the line segment L1 orthogonal to the velocity vector V1 at the tip of the velocity vector V1 and the line segment L2 orthogonal to the velocity vector V2 at the tip of the velocity vector V2. be able to.

  This is because when the actual velocity vector is V (Vx, Vy) and the Doppler velocity in each direction observed is Vd (Vxd, Vyd), the relationship between the two is generally Vd = A as coordinate transformation of the oblique coordinate system. XV can be expressed. Here, A is a quadratic matrix determined by the transmission angle of the Doppler beam. Therefore, if Vd is actually observed, V can be calculated as V = A−1 × Vd. However, A-1 is an inverse matrix of A.

  The blood flow velocity vector obtained in this way is stored in the image memory unit 10 and then averaged by the average vector calculator 11 to calculate an average flow velocity vector. Further, the average vector calculator 11 generates graphic data indicating the blood flow based on the calculated average flow velocity vector. The average vector calculator 11 can also use the graphic data as data indicating a graphic indicating the measurement region of the blood flow velocity vector.

  The graphic data generated by the average vector calculator 11 is given to the frame synthesizer 15 as frame data via the data synthesizer 14. This frame data is pixel-synthesized by a frame synthesizer 15 into a frame of image data reconstructed from a signal output from the B-mode processing circuit 4 or the TDI mode processing circuit 6 in the DSC 8. As a result, a display image obtained by combining the reconstructed image with the graphic is generated, and this display image is displayed on the display unit 16.

FIG. 5 is a diagram showing an example of a display image in which a graphic indicating the blood flow and the blood flow velocity vector measurement region is combined with the B-mode reconstructed image.
A graphic G1 shown in FIG. 5 indicates whether the blood flow is in the left-right direction in the image by hue and the flow velocity by lightness. The graphic G2 shows the outer edge of the blood flow velocity vector measurement region.

  Now, based on the blood flow velocity vector calculated as described above, a feature amount such as a blood flow rate can be calculated. In calculating the blood flow, ROI is set for that purpose. First, it is assumed that ROI R1 having a center position in the graphic G1 is designated by a user operation on a console (not shown) as shown in FIG. Then, the ROI generator 12 automatically sets the ROI R2 tilted so as to be orthogonal to the blood flow velocity vector while keeping the center position and shape unchanged. In the case of a rectangular ROI as shown in FIG. 6, for example, the long side is orthogonal to the blood flow velocity vector.

  The feature amount calculator 13 acquires information on the range of ROI R2 from the ROI generator 12, and calculates the blood flow rate based on the blood flow velocity vector calculated for the range of this ROI R2. The blood flow rate can be obtained by obtaining a velocity distribution on ROI R2 and rotationally integrating it with respect to the central axis of the blood vessel. That is, if the flow velocity at each point is Vi and the minute area of the semi-ring is dA, the blood flow rate Q is obtained as Q = Σvi × dA.

Next, attention is paid when the angle marker is moved on the display image by a user operation on a console (not shown). If the outside of the graphic G1 is designated as the position of the angle marker, the ROI generator 12 outputs graphic data for displaying the angle marker A1 having a default inclination at the designated position as shown in FIG. However, if the inside of the graphic G1 is designated as the position of the angle marker, as shown in FIG. 2, the ROI generator 12 designates the angle marker A2 in which the pointing direction is matched with the average flow velocity vector at the designated position. Output graphic data to be displayed on the screen.

  As described above, according to the present embodiment, the blood flow velocity vector is calculated based on the velocity vector observed two by two for one measurement site by Doppler beam scanning from two different directions. realizable.

  Further, according to the present embodiment, since the crossing angle between Doppler beams in different Doppler beam scans is always constant, the inverse matrix vector is constant at any measurement site, and the process of calculating the blood flow velocity vector is simplified.

  Further, according to the present embodiment, the graphic G1 indicating the measurement region of the blood flow velocity vector is displayed, so which region of the display image the blood flow velocity vector calculated as described above is reflected on. Can be easily recognized by an observer of the display image.

  According to the present embodiment, since the ROI inclination is automatically set so that the ROI for calculating the characteristic amount related to the blood flow such as the blood flow is orthogonal to the blood flow, the operator can set the ROI. There is no need to manually adjust the tilt. This reduces the burden on the operator and improves the accuracy and reproducibility of ROI settings.

  Further, according to the present embodiment, when the position of the angle marker is specified in the blood flow, the orientation direction of the angle marker is matched with the average flow velocity vector at that position, so that the blood more detailed than the direction that can be identified from the color display. The flow direction can be identified.

This embodiment can be variously modified as follows.
The direction of the Doppler beam may be changed in each Doppler beam scan. In this way, it is possible to enlarge a region where two Doppler beam scans overlap, that is, a region where a blood flow velocity vector can be measured. However, since the crossing angle between Doppler beams in different Doppler beam scans changes for each measurement site, it is necessary to change the inverse vector matrix for each measurement site, and the process of calculating the blood flow velocity vector becomes complicated.

  In the above embodiment, two Doppler beam scans are performed to calculate the blood flow velocity vector. Therefore, if this is performed for each frame, the frame rate is greatly reduced. When the frame rate is 20 frames / second, it is generally considered that there is no need to measure the blood flow velocity vector every frame, that is, every 50 ms. If the frame rate is more important than the real-time property of the blood flow velocity vector, the blood flow velocity vector is measured at a rate of one frame among a plurality of frames, and the blood flow velocity vector is calculated in the other frames. These two Doppler beam scans may not be performed. The frequency at which the blood flow velocity vector is measured may be fixed in advance, but if it can be changed according to the request from the operator, convenience is improved.

  In the above-described embodiment, the velocity vector of blood flow is calculated, and the moving body for which the velocity vector is calculated is blood. However, the velocity vector calculation target may be another moving body such as a tissue.

  Further, the function of automatically setting the inclination of the ROI is effective not only for analyzing blood flow but also for analyzing tissue motion, particularly heart wall motion. In tissue motion analysis, a rectangular or elliptical ROI is usually used, and an average velocity or a temporal change in velocity in the ROI is often measured. In this case, conventionally, the direction of movement is not known, or an operation such as manually assuming the direction of movement is necessary. Therefore, the ROI direction is not particularly taken into consideration, for example, ROI R11 and R12 shown in FIG. The horizontal direction on the screen was set as the axis. In order to adjust the ROI to the direction of the tissue motion, it was necessary to manually correct the inclination of the ROI. Even if the ROI inclination is set appropriately manually in the initial state, the ROI inclination may become inappropriate over time if the motion direction slightly changes over time. For this reason, it has been difficult to accurately measure a feature amount in a specific local region, for example, the endocardial side or the epicardial side. However, by applying the function of automatically setting the inclination of the ROI as shown in the above embodiment, the long side of the rectangular ROI with respect to the direction of tissue movement, for example, ROI R13 and R14 shown in FIG. Alternatively, it is possible to set an ROI in which the major axes of the ellipse ROI are orthogonal. FIG. 8 is a schematic diagram of a short-axis image of the heart, and shows a state where ROI R13 and R14 are placed on the myocardium. In the case of a tissue such as the heart, the target tissue moves, so the position of the ROI also needs to be moved. This can be applied in combination with a tracking method proposed by the prior art. . Thereby, an appropriate ROI corresponding to the actual tissue motion is set in each frame, and accurate feature amount calculation can be easily performed from the information in the appropriate ROI.

  Information on the velocity vector used to realize the function of automatically setting the tilt of the ROI and the function of automatically setting the pointing direction of the angle marker is not limited to that obtained by the method shown in the above embodiment, It is also possible to use another conventional method, for example, a method obtained based on the direction of movement of a living body obtained in advance by a technique such as automatic contour extraction or pattern matching. In this case, the present invention can be applied to another type of medical image analysis apparatus that calculates a velocity vector of a moving body without using ultrasonic waves.

  In the above embodiment, a normal two-dimensional ultrasonic diagnostic apparatus has been described in detail, but in recent years, an ultrasonic diagnostic apparatus capable of collecting a three-dimensional image in real time has been proposed. Applicable. In this case, an arbitrary cross-section is cut out to form a two-dimensional image, and the same processing as in the above embodiment may be applied thereto, or transmission / reception from three-dimensionally different directions may be performed. A motion speed vector may be obtained.

  Note that the present invention is not limited to the above-described embodiment as it is, and can be embodied by modifying the constituent elements without departing from the scope of the invention in the implementation stage. In addition, various inventions can be formed by appropriately combining a plurality of components disclosed in the embodiment. For example, some components may be deleted from all the components shown in the embodiment.

1 is a diagram showing a configuration of an ultrasonic diagnostic apparatus according to an embodiment of the present invention. The figure which shows the mode of two Doppler beam scans for calculating a blood-flow velocity vector. The figure which shows the measurement area | region of a blood-flow velocity vector. The figure which shows the principle which calculates a blood-flow velocity vector. The figure which shows an example of the display image which synthesize | combined the graphic which showed the measurement area | region of the blood flow and the blood flow velocity vector to the reconstruction image of B mode. The figure which shows a mode that the inclination of ROI is set automatically. The figure which shows a mode that the orientation direction of an angle marker is set automatically. It is a schematic diagram of a short-axis image of the heart, and shows a state in which the inclination of the ROI placed on the myocardium is automatically set.

Explanation of symbols

  DESCRIPTION OF SYMBOLS 1 ... Ultrasonic probe, 2 ... Transmission system circuit, 3 ... Reception system circuit, 4 ... B mode processing circuit, 5 ... CFM mode processing circuit, 6 ... TDI mode processing circuit, 7 ... PWD mode processing circuit, 8 ... Digital scan Converter (DSC) 9 ... Speed vector calculator 10 ... Image memory unit 11 ... Average vector calculator 12 ... ROI generator 13 ... Feature quantity calculator 14 ... Data synthesizer 15 ... Frame synthesizer 16 ... display unit, A1, A2 ... angle marker, G1, G2 ... graphic, P ... subject.

Claims (10)

Transmitting and receiving means for transmitting and receiving a Doppler beam at two different angles with respect to a plurality of measurement sites within a measurement range;
Based on two velocity vectors obtained individually from two Doppler beams received at each of the two angles for one measurement portion of the plurality of measurement portions, the velocity vector of the moving body at the one measurement portion is calculated. Calculating means for executing processing for each of the plurality of measurement sites;
Generating means for generating an image representing a motion state of the moving body within the measurement range based on the plurality of velocity vectors respectively calculated by the calculating means with respect to the plurality of measurement sites ;
Setting means for setting a rectangular region of interest in response to a user operation in the image;
Changing means for changing the direction of the region of interest so that one side of the region of interest is orthogonal to the velocity vector of the moving body in the region of interest set by the setting unit;
An ultrasonic diagnostic apparatus comprising: means for calculating a feature quantity of the moving body in the region of interest whose direction has been changed .   The said calculation means calculates the velocity vector of the said moving body by performing coordinate transformation of a skew coordinate system with respect to two velocity vectors calculated | required separately from the said two Doppler beams. Ultrasound diagnostic equipment.   The ultrasonic diagnostic apparatus according to claim 1, wherein the generation unit generates the image based on a vector obtained by spatially averaging the plurality of velocity vectors calculated by the calculation unit.   The ultrasonic diagnostic apparatus according to claim 1, wherein the transmitting / receiving unit transmits and receives the two Doppler beams so that an angle formed by the two Doppler beams is constant in any of the plurality of measurement sites.   2. The transmission / reception unit transmits and receives the two Doppler beams so as to change an angle formed by the two Doppler beams with respect to other measurement sites in at least a part of the plurality of measurement sites. The ultrasonic diagnostic apparatus as described.   The ultrasonic diagnostic apparatus according to claim 1, wherein the generation unit generates a color Doppler image representing a motion state of the moving body.   The ultrasonic diagnostic apparatus according to claim 1, wherein the generation unit generates an image that further represents the measurement range. The ultrasonic diagnostic apparatus according to claim 1 , wherein the changing unit changes the direction of the region of interest for each frame based on the images of a plurality of frames in time series. The ultrasonic diagnostic apparatus according to claim 1 , wherein the generation unit generates an image that further represents an average velocity vector in the region of interest.   The ultrasonic wave according to claim 1, further comprising means for adding an image representing a velocity vector detected by the detecting means with respect to a designated position in the image to the image generated by the generating means. Diagnostic device.

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