WO2022181022A1 - Dispositif et procédé de traitement d'image, système de radiographie et programme - Google Patents

Dispositif et procédé de traitement d'image, système de radiographie et programme Download PDF

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WO2022181022A1
WO2022181022A1 PCT/JP2021/047539 JP2021047539W WO2022181022A1 WO 2022181022 A1 WO2022181022 A1 WO 2022181022A1 JP 2021047539 W JP2021047539 W JP 2021047539W WO 2022181022 A1 WO2022181022 A1 WO 2022181022A1
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image
images
moving
frame
energy
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Japanese (ja)
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竜一 藤本
貴司 岩下
晃介 照井
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キヤノン株式会社
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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment

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  • the present invention relates to an image processing apparatus and method, a radiation imaging system, and a program.
  • radiographic imaging devices that use flat panel detectors (Flat Panel Detectors, hereinafter abbreviated as FPDs) made of semiconductor materials are widely used as imaging devices for X-ray medical image diagnosis and non-destructive inspection.
  • FPDs Flat Panel detectors
  • Such radiation imaging apparatuses are used, for example, in medical image diagnosis as digital imaging apparatuses for still image capturing such as general radiography and moving image capturing such as fluoroscopic imaging.
  • One of imaging methods using an FPD is energy subtraction (Patent Document 1).
  • a plurality of images corresponding to X-rays of different energies are acquired, and an image of a specific material (for example, a bone image and a soft tissue image) is obtained from the plurality of images by utilizing the difference in the X-ray attenuation rate of the material. tissue images) are separated.
  • a specific material for example, a bone image and a soft tissue image
  • Patent Document 2 describes a system that performs dynamic dual-energy imaging by applying energy subtraction processing to moving images.
  • the tube voltage of the X-ray source is changed to the first kV value and then to the second kV value.
  • the first signal corresponding to the first sub-image is integrated when the tube voltage is at the first kV value, and the integration is reset after the integrated signal is transferred to the sample and hold node.
  • a second signal corresponding to a second sub-image is then integrated when the tube voltage is at a second kV value.
  • readout of the integrated first signal and integration of the second signal are performed in parallel.
  • an energy subtraction moving image can be captured.
  • the above-described conventional technology has a problem that flickering occurs in the moving image because the pixel values of the energy subtraction moving image vary between frames due to temporal output fluctuations between sub-image frames. Energy-subtracted movies with flickering may lead to inaccurate diagnostic imaging.
  • the present invention provides a technique for reducing flickering in energy subtraction moving images.
  • An image processing apparatus has the following configuration. i.e. generating means for generating moving images of separated images by performing energy subtraction processing using a plurality of moving images corresponding to a plurality of different radiation energies; and correction means for correcting the moving images of the plurality of moving images or the moving images of the separated images so as to reduce variations in signal values between frames in the moving images of the plurality of moving images or the moving images of the separated images.
  • FIG. 1 is a diagram showing a configuration example of a radiation imaging system according to an embodiment
  • FIG. FIG. 2 is an equivalent circuit diagram of pixels included in a two-dimensional detector of the X-ray imaging apparatus
  • 4 is a timing chart showing operations for acquiring an X-ray image
  • 4 is a timing chart showing operations for acquiring an X-ray image
  • FIG. 4 is a block diagram of correction processing in energy subtraction processing
  • FIG. 3 is a block diagram of image processing related to energy subtraction processing;
  • FIG. 2 is a block diagram showing a hardware configuration example of a control computer
  • FIG. 4 is a diagram showing an example of accumulated images and bone images
  • FIG. 4 is a diagram showing signal value fluctuations between frames of a moving image before and after energy subtraction processing
  • 4 is a block diagram of signal processing including blink reduction processing according to the first embodiment
  • FIG. 4 is a flowchart of flickering reduction processing according to the embodiment
  • FIG. 4 is a diagram showing signal value fluctuations between frames of a moving image subjected to flickering reduction processing
  • 4A and 4B are diagrams for explaining a process of acquiring an accumulated image A
  • FIG. FIG. 11 is a block diagram of signal processing including flicker reduction processing and noise reduction processing according to the second embodiment
  • Radiation in the present invention includes alpha rays, beta rays, and gamma rays, which are beams produced by particles (including photons) emitted by radioactive decay, as well as beams having radiation energy at the same level or higher, such as particle beams. , cosmic rays, etc. are also included.
  • FIG. 1 is a block diagram showing a configuration example of a radiation imaging system according to the first embodiment.
  • the radiation imaging system of the first embodiment includes an X-ray generation device 101 , an X-ray control device 102 , a control computer 103 and an X-ray imaging device 104 .
  • the X-ray generator 101 emits X-rays.
  • the X-ray controller 102 controls X-ray irradiation by the X-ray generator 101 .
  • the control computer 103 controls the X-ray imaging device 104 to acquire a radiographic image (hereinafter referred to as an X-ray image (image information)) captured by the X-ray imaging device 104 .
  • the control computer 103 functions as an image processing device that performs image processing including flickering reduction processing, which will be described later, on an X-ray image acquired from the X-ray imaging device 104 .
  • the X-ray imaging apparatus 104 may be provided with a function of executing image processing including flickering reduction processing (and noise reduction processing described in the second embodiment).
  • the X-ray imaging device 104 is composed of a phosphor 105 that converts X-rays into visible light and a two-dimensional detector 106 that detects visible light.
  • the two-dimensional detector 106 is a sensor in which pixels 20 for detecting X-ray quanta are arranged in an array of X columns ⁇ Y rows, and outputs image information.
  • FIG. 5 is a block diagram showing a hardware configuration example of the control computer 103.
  • the CPU 141 controls various operations of the control computer 103 by executing programs stored in the ROM 142 or RAM 143 .
  • the CPU 141 controls X-ray irradiation by the X-ray control device 102 (X-ray generator 101 ) and X-ray image capturing operation by the X-ray imaging device 104 .
  • the CPU 141 also implements various signal processing and image processing, which will be described later. It should be noted that the operation of signal processing and image processing, which will be described later, may be partially or wholly realized by dedicated hardware.
  • the ROM 142 stores programs executed by the CPU 141 and various data.
  • the RAM 143 provides a work area for storing intermediate data generated when the CPU 141 executes processing.
  • the secondary storage device 144 stores radiation images (X-ray images) to be processed.
  • the secondary storage device 144 also stores control programs.
  • the programs stored in the secondary storage device 144 are developed in the RAM 143 as necessary and executed by the CPU 141 .
  • the display 145 performs various displays under the control of the CPU 141.
  • the operating unit 146 includes, for example, a keyboard and a pointing device, and receives various user inputs.
  • the interface 147 connects external devices such as the X-ray control device 102 and the X-ray imaging device 104 to the control computer 103 .
  • a bus 148 communicably connects the units described above.
  • FIG. 2 is an equivalent circuit diagram of the pixel 20 included in the two-dimensional detector 106.
  • the pixel 20 includes a photoelectric conversion element 201 and an output circuit section 202 .
  • Photoelectric conversion element 201 can typically be a photodiode.
  • the output circuit section 202 includes an amplifier circuit section 204 , a clamp circuit section 206 , a sample hold circuit section 207 and a selection circuit section 208 .
  • the photoelectric conversion element 201 includes a charge storage section, and the charge storage section is connected to the gate of the MOS transistor 204 a of the amplifier circuit section 204 .
  • the source of MOS transistor 204a is connected to current source 204c through MOS transistor 204b.
  • a source follower circuit is formed by the MOS transistor 204a and the current source 204c.
  • the MOS transistor 204b is an enable switch that turns on when the enable signal EN supplied to its gate becomes active level to put the source follower circuit into operation.
  • the charge-voltage converter is connected to reset potential Vres through reset switch 203 . When the reset signal PRES becomes active level, the reset switch 203 is turned on, and the potential of the charge-voltage converter is reset to the reset potential Vres.
  • the clamp circuit section 206 clamps the noise output by the amplifier circuit section 204 according to the reset potential of the charge-voltage conversion section with the clamp capacitor 206a.
  • the clamp circuit unit 206 is a circuit for canceling this noise from the signal output from the source follower circuit according to the charge generated by photoelectric conversion in the photoelectric conversion element 201 .
  • This noise includes kTC noise at reset. Clamping is performed by setting the clamp signal PCL to the active level to turn on the MOS transistor 206b and then setting the clamp signal PCL to the inactive level to turn off the MOS transistor 206b.
  • the output side of the clamp capacitor 206a is connected to the gate of the MOS transistor 206c.
  • MOS transistor 206c The source of MOS transistor 206c is connected to current source 206e through MOS transistor 206d.
  • a source follower circuit is formed by the MOS transistor 206c and the current source 206e.
  • the MOS transistor 206d is an enable switch that turns on when the enable signal EN0 supplied to its gate becomes active level to put the source follower circuit into operation.
  • a signal output from the clamp circuit unit 206 according to the charge generated by photoelectric conversion in the photoelectric conversion element 201 is written as a light signal into the capacitor 207Sb via the switch 207Sa when the light signal sampling signal TS becomes active level.
  • the signal output from the clamp circuit section 206 when the MOS transistor 206b is turned on immediately after resetting the potential of the charge-voltage conversion section is the clamp voltage.
  • This noise signal is written into the capacitor 207Nb through the switch 207Na when the noise sampling signal TN becomes active level.
  • This noise signal contains the offset component of the clamp circuit section 206 .
  • a switch 207Sa and a capacitor 207Sb constitute a signal sample and hold circuit 207S
  • a switch 207Na and a capacitor 207Nb constitute a noise sample and hold circuit 207N.
  • the sample and hold circuit section 207 includes a signal sample and hold circuit 207S and a noise sample and hold circuit 207N.
  • the drive circuit drives the row selection signal to the active level
  • the signal (light signal) held in the capacitor 207Sb is output to the signal line 21S via the MOS transistor 208Sa and the row selection switch 208Sb.
  • the signal (noise) held in capacitor 207Nb is output to signal line 21N via MOS transistor 208Na and row select switch 208Nb.
  • the MOS transistor 208Sa forms a source follower circuit with a constant current source (not shown) provided on the signal line 21S.
  • the MOS transistor 208Na forms a source follower circuit with a constant current source (not shown) provided on the signal line 21N.
  • a signal selection circuit portion 208S is composed of the MOS transistor 208Sa and the row selection switch 208Sb
  • a noise selection circuit portion 208N is composed of the MOS transistor 208Na and the row selection switch 208Nb.
  • the selection circuit section 208 includes a signal selection circuit section 208S and a noise selection circuit section 208N.
  • the pixel 20 may have an addition switch 209S that adds the optical signals of a plurality of adjacent pixels 20.
  • the addition mode signal ADD becomes active level and the addition switch 209S is turned on.
  • the capacitors 207Sb of adjacent pixels 20 are connected to each other by the addition switch 209S, and the optical signals are averaged.
  • pixel 20 may have a summing switch 209N that sums the noise of adjacent pixels 20 .
  • Addition section 209 includes an addition switch 209S and an addition switch 209N.
  • the pixel 20 may have a sensitivity changing section 205 for changing sensitivity.
  • the pixel 20 can include, for example, a first sensitivity change switch 205a and a second sensitivity change switch 205'a and their associated circuit elements.
  • the first change signal WIDE becomes active level
  • the first sensitivity change switch 205a is turned on, and the capacitance value of the first additional capacitor 205b is added to the capacitance value of the charge-voltage converter. This reduces the sensitivity of the pixel 20 .
  • the second change signal WIDE2 becomes active level
  • the second sensitivity change switch 205'a is turned on, and the capacitance value of the second additional capacitor 205'b is added to the capacitance value of the charge-voltage converter.
  • the enable signal ENw may be made the active level to cause the MOS transistor 204'a to perform the source follower operation instead of the MOS transistor 204a.
  • the X-ray imaging apparatus 104 reads the output of the pixel circuit as described above, converts it into a digital value with an AD converter (not shown), and then transfers the image to the control computer 103 .
  • 3A and 3B are diagrams showing drive timings of the X-ray imaging apparatus 104 according to the first embodiment.
  • the horizontal axis represents time, and the timings of X-ray irradiation, synchronization signals, resetting of the photoelectric conversion element 201 , sample hold circuit 207 and image reading from the signal line 21 are shown.
  • the photoelectric conversion element 201 is reset, and then X-rays are emitted.
  • the X-ray tube voltage ideally becomes a rectangular wave, it takes a finite amount of time for the tube voltage to rise and fall.
  • the tube voltage can no longer be regarded as a rectangular wave, and has a waveform as shown in FIG. 3A. That is, the radiation energy (energy of X-rays) differs in the rising period, the stable period, and the falling period of X-rays.
  • sampling is performed by the noise sample-and-hold circuit 207N after the X-ray 301 in the rising period is emitted, and sampling is performed by the signal sample-and-hold circuit 207S after the X-ray 302 in the stable period is emitted.
  • the difference between the signal lines 21N and 21S is read out as an image.
  • the noise sample-and-hold circuit 207N holds the signal (R 1 ) of the X-rays 301 in the rising period
  • the signal sample-and-hold circuit 207S holds the signal of the X-rays 301 in the rising period and the signal of the X-rays 302 in the stable period. (R 1 +B) is retained. Therefore, an image 304 corresponding to the signal (B) of the X-rays 302 in the stable period is read out from the X-ray imaging apparatus 104 .
  • sampling is performed again by the signal sample-and-hold circuit 207S.
  • the photoelectric conversion element 201 is reset, sampling is performed again by the noise sample hold circuit 207N, and the difference between the signal lines 21N and 21S is read out as an image.
  • the noise sample-and-hold circuit 207N holds the signal in the state where the X-ray is not irradiated.
  • the signal sample-and-hold circuit 207S holds the sum (R 1 +B+R 2 ) of the signal of the X-ray 301 in the rising period, the X-ray 302 in the stable period, and the signal of the X-ray 303 in the falling period. Therefore, from the X-ray imaging apparatus 104, an image 306 corresponding to the signal of the X-rays 301 in the rise period, the signal of the X-rays 302 in the stable period, and the signal of the X-rays 303 in the fall period is read.
  • the timing for resetting the sample-and-hold circuit 207 and the photoelectric conversion element 201 is determined using the synchronization signal 307 indicating that the X-ray generator 101 has started to emit X-rays.
  • a configuration that measures the tube current of the X-ray generator 101 and determines whether or not the current value exceeds a preset threshold value is preferably used.
  • the photoelectric conversion element 201 is completely reset, the pixel 20 is repeatedly read out, and a configuration in which it is determined whether or not the pixel value exceeds a preset threshold value is preferably used.
  • a configuration in which an X-ray detector different from the two-dimensional detector 106 is incorporated in the X-ray imaging apparatus 104 and whether or not the measured value exceeds a preset threshold is preferably used.
  • sampling of the signal sample and hold circuit 207S, sampling of the noise sample and hold circuit 207N, and resetting of the photoelectric conversion element 201 are performed.
  • an image 304 corresponding to the stable period of the pulse X-ray and an image 305 corresponding to the sum of the rising period and the falling period are obtained. Since the energies of the X-rays irradiated when forming the two images are different, energy subtraction processing can be performed by performing calculations between the images.
  • FIG. 3B shows drive timing when energy subtraction is performed in the radiation imaging system according to the first embodiment. This differs from FIG. 3A in that the X-ray tube voltage is actively switched.
  • the photoelectric conversion element 201 is reset, and then low-energy X-rays 401 are irradiated. After that, sampling is performed by the noise sample-and-hold circuit 207N, and after the tube voltage is switched and the high-energy X-ray 402 is emitted, sampling is performed by the signal sample-and-hold circuit 207S. Thereafter, the tube voltage is switched to irradiate low-energy X-rays 403 . Furthermore, the difference between the signal lines 21N and 21S is read out as an image.
  • the noise sample-and-hold circuit 207N holds the signal (R 1 ) of the low-energy X-rays 401
  • the signal sample-and-hold circuit 207S holds the signal of the low-energy X-rays 401 and the signal of the high-energy X-rays 402. (R 1 +B) is retained. Therefore, an image 404 corresponding to the signal (B) of high-energy X-rays 402 is read out from the X-ray imaging apparatus 104 .
  • sampling is performed again by the signal sample-and-hold circuit 207S.
  • the photoelectric conversion element 201 is reset, sampling is performed again by the noise sample hold circuit 207N, and the difference between the signal lines 21N and 21S is read out as an image.
  • the noise sample-and-hold circuit 207N holds the signal in the state where the X-ray is not irradiated.
  • the sync signal 407 is the same as in FIG. 3A.
  • the energy subtraction processing in this embodiment is divided into three stages of correction processing, signal processing, and image processing. Processing at each stage will be described below.
  • FIG. 4A is a block diagram of correction processing in the energy subtraction processing according to this embodiment.
  • imaging is performed without exposing the X-ray imaging device 104 to X-rays, and an image is acquired by the driving shown in FIG. 3A or 3B.
  • two images are read, the first image being F_ODD and the second image being F_EVEN.
  • F_ODD and F_EVEN are images corresponding to fixed pattern noise (FPN) of the X-ray imaging device 104 .
  • FPN fixed pattern noise
  • W_ODD and W_EVEN are images corresponding to the sum of signals from the FPN of the X-ray imaging device 104 and X-rays. Therefore, by subtracting F_ODD from W_ODD and F_EVEN from W_EVEN, the FPN-removed images WF_ODD and WF_EVEN of the X-ray imaging apparatus 104 are obtained. This is called offset correction.
  • WF_ODD is an image corresponding to X-rays 302 (or high-energy X-rays 402) in the stable period.
  • WF_EVEN is the sum of rising phase X-ray 301, stable phase X-ray 302, and falling phase X-ray 303 (or the sum of low energy X-rays 401 and 403 and high energy X-ray 402). is an image corresponding to Therefore, by subtracting WF_ODD from WF_EVEN, an image corresponding to the sum of the X-rays 301 in the rising period and the X-rays 303 in the falling period is obtained.
  • the energies of the X-rays 301 in the rising period and the X-rays 303 in the falling period are lower than the energy of the X-rays 302 in the stable period. Therefore, by subtracting WF_ODD from WF_EVEN, a low energy image W_Low without a subject is obtained. Also, from WF_ODD, a high-energy image W_High with no subject is obtained. This is called color correction.
  • an X-ray is emitted to the X-ray imaging device 104 in a state where an object is present to perform imaging, and an image is acquired by the driving shown in FIG. 3A or 3B.
  • two images are read, the first image being X_ODD and the second image being X_EVEN.
  • d be the thickness of the subject
  • ⁇ be the linear attenuation coefficient of the subject
  • I 0 be the output of the pixel 20 when there is no subject
  • I be the output of the pixel 20 when there is the subject. holds.
  • Equation (2) By transforming the formula (1), the following formula (2) is obtained.
  • the right side of Equation (2) indicates the attenuation rate of the object.
  • the attenuation rate of the subject is a real number between 0 and 1.
  • FIG. 4B shows a block diagram of signal processing in energy subtraction processing.
  • separated images are obtained from the attenuation rate image L at low energy and the attenuation rate image H at high energy obtained by the correction processing shown in FIG. 4A.
  • a case of obtaining a bone thickness image B (also referred to as a bone image B) and a soft tissue thickness image S (also referred to as a soft tissue image S) will be described.
  • E the energy of X-ray photons
  • N(E) the number of photons at energy E
  • B the thickness of bone
  • S the thickness of soft tissue.
  • the number of photons N(E) at energy E is the X-ray spectrum.
  • the X-ray spectrum is obtained by simulation or actual measurement.
  • the linear attenuation coefficient ⁇ B (E) of bone at energy E and the linear attenuation coefficient ⁇ S (E) of soft tissue at energy E can be obtained from databases such as NIST. That is, it is possible to calculate the attenuation rate I/I 0 for any bone thickness B, soft tissue thickness S, X-ray spectrum N(E).
  • the thickness B of the bone and the thickness S of the soft tissue can be obtained.
  • a case of using the Newton-Raphson method as a representative method for solving nonlinear simultaneous equations will be described.
  • m be the number of iterations of the Newton-Raphson method
  • Bm be the thickness of the bone after the mth iteration
  • Sm be the thickness of the soft tissue after the mth iteration.
  • the attenuation rate is H m
  • the low-energy attenuation rate L m after the m-th iteration is represented by the following equation (5).
  • the change rate of the attenuation rate when the thickness changes minutely is expressed by the following formula (6).
  • the bone thickness B m+1 and the soft tissue thickness S m+1 after the m+1th iteration are represented by the following equation (7) using the high-energy attenuation rate H m and the low-energy attenuation rate L m . .
  • the inverse matrix of a 2 ⁇ 2 matrix is represented by the following formula (8) from Cramer's formula, where det is the determinant.
  • the difference between the high-energy attenuation rate Hm after the m -th iteration and the actually measured high-energy attenuation rate H approaches zero limitlessly.
  • the same is true for the attenuation rate L of low energy.
  • the bone thickness Bm after the mth iteration converges to the bone thickness B
  • the soft tissue thickness Sm converges to the soft tissue thickness S after the mth iteration.
  • the nonlinear simultaneous equations shown in Equation (4) can be solved. Therefore, by calculating Equation (4) for all pixels, the bone thickness image B and the soft tissue thickness image B are obtained from the attenuation rate image L at low energy and the attenuation rate image H at high energy. S can be obtained.
  • the embodiment is not limited to such a form.
  • the thickness W of water and the thickness I of the contrast agent may be calculated as separate images. That is, it may be decomposed into thicknesses of any two types of materials.
  • an image of the effective atomic number Z and an image of the surface density D may be obtained as separate images from the image L of the attenuation rate at low energy and the image H of the attenuation rate at high energy obtained by the correction process of FIG. .
  • the effective atomic number Z is the equivalent atomic number of a mixture
  • the areal density D is the product of the object density [g/cm 3 ] and the object thickness [cm].
  • the signal processing of the present invention can be said to be processing for generating an energy subtraction image by computing a low energy image and a high energy image (energy subtraction processing).
  • energy subtraction image and separation image are synonymous.
  • an example of solving the nonlinear simultaneous equations using the Newton-Raphson method is shown, but the method is not limited to such a form.
  • an iterative solution method such as the least squares method or the bisection method may be used.
  • the nonlinear simultaneous equations are solved by the iterative solution method, but the present invention is not limited to such a form.
  • Bone thickness B and soft tissue thickness S for various combinations of high energy attenuation rate H and low energy attenuation rate L are obtained in advance to generate a table, and by referring to this table, bone thickness B and A configuration for obtaining the thickness S of the soft tissue at high speed may be used.
  • FIG. 4C shows a block diagram of image processing related to energy subtraction processing.
  • an image for display is generated using the separated images obtained by the signal processing described above.
  • a display image is generated by post-processing the bone image B obtained by the signal processing shown in FIG. 4B.
  • the generated display image is displayed on the display 145, for example.
  • Such post-processing may include logarithmic transformation, dynamic range compression, and the like.
  • the content of processing may be switched by inputting the type and strength of post-processing as parameters.
  • FIG. 6 schematically shows an example of an accumulated image 601 and a bone image 602.
  • the stored image 601 is an image before energy subtraction processing, that is, an image captured by an existing radiation imaging system without energy resolution or an image corresponding thereto.
  • image 306, image 406 in FIGS. 3A and 3B, high energy image H, and low energy image L described above correspond to accumulated images.
  • Bone image 602 is a separated image obtained by the energy subtraction process described above.
  • the normal human body consists only of soft tissue and bone.
  • Interventional Radiology hereinafter referred to as IVR
  • a contrast agent is injected into the blood vessel.
  • Treatments such as inserting a catheter or guide wire into a blood vessel and placing a stent or coil are also performed.
  • IVR IVR
  • treatment is performed while confirming the positions and shapes of contrast agents and medical devices. Therefore, it is desired to improve the visibility by isolating only the contrast agent and medical device or by removing the background such as soft tissue and bone.
  • the soft tissue is visible, whereas in the bone image 602 obtained by the energy subtraction processing, the contrast of the soft tissue can be removed.
  • the main component of contrast media is iodine
  • the main component of medical devices is metal such as stainless steel. Since both have atomic numbers greater than calcium, which is the main component of bone, the bone image 602 displays the bone, the contrast medium, and the medical device.
  • the same effect can be obtained by changing the tube voltage and filter for low-energy X-rays and high-energy X-rays. In either case, it was confirmed that the bone image 602 displayed the bone, the contrast agent, and the medical device.
  • contrast enhancement It may improve the visibility of drugs and medical devices.
  • the bone image 602 has a larger frame-to-frame variation than the accumulated image 601, and there is a problem that the screen flickers when a moving image is displayed.
  • the pixel value of the accumulated image 601 varies for each shooting (frame). Possible causes include X-ray source-related factors such as variations in exposure dose and radiation quality between frames, and sensor-related factors such as sample-and-hold timing changes between frames. Variations in pixel values in an accumulated image are amplified by performing energy subtraction processing.
  • FIG. 7 shows inter-frame variation with respect to the average image of frames before and after the energy subtraction process when radiographic moving images of porcine organs are captured.
  • a graph 701 shows inter-frame variation in a moving image (bone image) after energy subtraction processing.
  • a graph 702 represents inter-frame variation in the moving image (high energy image H) before the energy subtraction process. However, the image before the energy subtraction process adjusts the pixel value to [cm] by the following deformation.
  • each pixel I/ I0 of the image before energy subtraction can be expressed by equation (2) .
  • ⁇ is the attenuation rate at the average energy E of X-rays.
  • equation (2) can be transformed into equation (10) below.
  • each pixel I/ I0 of the image before the energy subtraction process as shown in Equation (10), it can be made to have the dimension of thickness [cm].
  • the bone attenuation rate is used as the attenuation rate ⁇ .
  • FIG. 8 shows a block diagram of signal processing including processing for reducing fluctuations (blinking) in signal values between frames (hereinafter, flickering reduction processing).
  • the block diagram of FIG. 8 shows in more detail the signal processing blocks described above with reference to FIG. 4B.
  • separated images bone image B′ and soft tissue A moving image of the tissue image S'
  • flickering between frames is reduced by flickering reduction processing, which will be described later.
  • FIG. 8 shows the minimum required configuration for the sake of explanation.
  • block R1, block MD1, and block R2 can be realized by executing a program stored in ROM 142 by CPU 141, for example.
  • any one or all of the blocks may be implemented by dedicated hardware, or may be implemented by cooperation between the CPU and dedicated hardware.
  • Block R1 performs flicker reduction processing on the high energy image H and the low energy image L. Details of the flickering reduction process will be described later.
  • Block MD1 then applies the signal processing described with reference to FIG. to generate Block R2 then performs a blink reduction process on bone image B and soft tissue image S to generate bone image B' and soft tissue image S'.
  • moving images B′ and S′ which are energy subtraction processed moving images subjected to flickering reduction processing, are obtained.
  • FIG. 9 is a flowchart showing the flickering reduction process performed by block R1 shown in FIG.
  • block R1 calculates representative signal values from high energy image H and low energy image L, respectively.
  • the representative signal value is unaffected by subject motion.
  • a representative signal value may be obtained from the entire image, or may be obtained from a set ROI.
  • the ROI may be a predetermined fixed region, may be set manually by the user when calculating the representative signal value, or may be set automatically.
  • a method of automatically setting the ROI for example, a method of analyzing a moving image to specify an area where there is no signal value variation due to the object, and setting the specified area as the ROI can be used.
  • a statistic such as the median value, the average value, or the N% value of the cumulative histogram can be used as the representative value signal.
  • the block R1 obtains a reference representative signal value (hereinafter referred to as reference value) from the moving image data to be subjected to flickering reduction processing.
  • the reference value can be, for example, the median value of the first frame image of the moving image data. More specifically, the median value of the image of the first frame of the moving image data composed of the series of high energy images H is used as the reference value for the high energy image, and the median value of the image of the first frame of the moving image data composed of the series of low energy images L The median value of the image is determined as the reference value for the low energy image.
  • the block R1 calculates the correction amount for each of the high energy image H and the low energy image L.
  • the correction amount ⁇ of the high-energy image H and the correction of the high-energy image H using the correction amount ⁇ will be described below, but the low-energy image can also be corrected by similar processing.
  • the correction amount ⁇ is obtained by calculating how many times the reference value obtained from the moving image data of the high energy image in S902 is the median value obtained for the high energy image H in S901.
  • block R1 multiplies the high energy image H by the correction amount ⁇ obtained at S903 to obtain a high energy image H′. In this way, the frame-to-frame variation of the signal in the moving image composed of the high-energy images H is reduced.
  • Equation (11) is an example of blinking reduction processing using multiplication of a coefficient (correction amount ⁇ ).
  • the reference value was obtained from the first frame image, but it is not limited to this.
  • the reference value may be calculated using all or part of the frames captured at the time the frame to be corrected was obtained (that is, the frames prior to the frame to be corrected).
  • the reference value may be calculated from the latest average image of a predetermined number of frames at the time when the frame to be corrected is acquired.
  • the correction amount is calculated independently for the high energy image and the low energy image, but the present invention is not limited to this. You can do it.
  • the high-energy image H and the low energy image L may be corrected.
  • the correction amount is obtained by division, and the correction is performed by multiplication, but the present invention is not limited to this.
  • the correction amount may be derived and corrected by addition/subtraction.
  • Equation (14) shows the case of obtaining a high energy image H′ subjected to blink reduction processing from the high energy image H by addition/subtraction.
  • Equation (14) is an example of blink reduction processing using addition and subtraction of a coefficient (correction amount ⁇ ).
  • the flickering reduction processing for correcting the images before the energy subtraction processing has been described above.
  • the flickering reduction processing performed by the block R1 on images (high energy image H and low energy image L) corresponding to radiography with a plurality of energies before separation of correction targets has been described above.
  • a block R2 performs flickering reduction processing with the separated image as the target of correction.
  • a method similar to the above-described flickering reduction processing for example, the method represented by Equation (11) or Equation (14) can be used for the flickering reduction processing.
  • FIG. 10 shows the frame-to-frame variation of the image average value in the moving image obtained by the energy subtraction processing with and without the above-described flickering reduction processing.
  • a graph 701 like FIG. 7, shows the variation of the image average value in the bone image when the flicker reduction processing is not performed.
  • a graph 1001 shows the variation of the image average value in the bone image when the flicker reduction processing is performed.
  • the flickering reduction process used for block R1 and block R2 may be either a method of multiplying coefficients or a method of adding or subtracting coefficients.
  • blinking can be suppressed most effectively when multiplication correction is performed in blinking reduction processing for the high energy image H and the low energy image L.
  • flickering can be suppressed most effectively when the bone image B and the soft tissue image S are corrected by addition and subtraction in the flickering reduction process. Therefore, it is preferable to apply correction by multiplication of coefficients to the blink reduction processing of block R1 and correction by addition and subtraction of coefficients to the blink reduction processing of block R2.
  • the flicker reduction process is performed on both the image before the energy subtraction process and the image after the energy subtraction process, but the present invention is not limited to this.
  • At least one of block R1 and block R2 may perform the blink reduction process.
  • the inventors of the present application examined the block R1 and the block R2, it was found that blinking could be reduced more satisfactorily by using both the block R1 and the block R2.
  • the noise reduction processing of this embodiment effectively reduces noise using stored images that are compatible with images acquired by general radiography.
  • an accumulated image used for noise reduction processing according to this embodiment will be described.
  • high-resolution images are obtained from a plurality of radiographic images acquired by sample-holding at a plurality of timings including timings during X-ray irradiation and after the end of X-ray irradiation.
  • An energy image and a low energy image are generated.
  • an X-ray image (image 306 in FIG. 3A) acquired at the timing after the end of X-ray irradiation can be used as the accumulated image.
  • FIG. 11 is a block diagram showing an example of a configuration for implementing correction processing for acquiring an accumulated image according to the second embodiment.
  • the accumulated image A is generated, for example, by dividing the image XF_EVEN by the image WF_EVEN.
  • Image XF_EVEN and image WF_EVEN are as described in FIG. 4A. That is, the image XF_EVEN is an image corresponding to the sum of the X-rays 301 in the rising period, the X-rays 302 in the stable period, and the X-rays 303 in the falling period when there is an object.
  • the image WF_EVEN is an image corresponding to the sum of the X-rays 301 in the rising period, the X-rays 302 in the stable period, and the X-rays 303 in the falling period when there is no subject.
  • the stored image A may be generated by multiplying an image H with an attenuation rate in high energy (high energy image H) and an image L with an attenuation rate in low energy (low energy image L) and adding them together.
  • accumulated image A may be generated using equation (15). Note that in calculating the accumulated image A, one coefficient may be set to 0 and the other coefficient may be set to 1. In this case, the high energy image H or the low energy image L itself is used as the accumulated image A. That is, an image captured at substantially the same timing as the image to be subjected to energy subtraction processing and to which energy subtraction processing has not been applied can be used as accumulated image A for noise reduction processing.
  • FIG. 12 is a block diagram showing an example of a configuration for performing signal processing in energy subtraction processing according to the second embodiment.
  • the block diagram of FIG. 12 shows a more detailed configuration of the signal processing described with reference to FIG. 4B. Note that FIG. 12 shows the minimum required configuration for the sake of explanation.
  • Blocks R1-R3, blocks MD1-MD2, block ADD, and blocks F1-F3 can be implemented by the CPU 141 executing a program stored in the ROM 142, for example. However, any one or all of the blocks shown in FIG. 12 may be implemented by dedicated hardware, or may be implemented by cooperation between the CPU and dedicated hardware.
  • Block F2 and block F3 perform filtering for the purpose of noise reduction on the low energy image L and the high energy image H, respectively.
  • Generate image H' For filtering, for example, spatial filters such as Gaussian filters and median filters, structure-preserving spatial filters such as epsilon filters and Laplacian filters, and time filters such as recursive filters can be used.
  • Quantum noise of the X-rays is reduced by a noise reduction process that applies a filter before the block MD1 that performs two-material separation.
  • Block R1 performs a blink reduction process on the noise reduced low energy image L' and the noise reduced high energy image H' to produce a high energy image H'' and a low energy image L''.
  • a moving image composed of the high-energy image H'' and the low-energy image L'' is a moving image in which inter-frame fluctuations of signals are suppressed, both of which are obtained from the block R1.
  • a block MD1 generates a bone image B' and a soft tissue image S' from the high energy H'' and low energy image L'' that have undergone noise reduction processing and are corrected for inter-frame variation.
  • the operation of material separation by block MD1 is the same as in the first embodiment (block MD1 in FIG. 8).
  • a block R2 performs flicker reduction processing on the bone image B' and the soft tissue image S' generated in the block MD1, and outputs a bone image B'' and a soft tissue image S''.
  • the processing by block R2 is the same as in the first embodiment (block RD2 in FIG. 8).
  • a block ADD generates an image of the sum of the bone image B'' and the soft tissue image S'', and outputs it as a thickness image T'.
  • the thickness image T' which is the sum of the bone image B'' and the soft tissue image S'', has undergone noise reduction and blink reduction.
  • a block F1 applies filter processing for the purpose of noise reduction to the thickness image T' to generate a thickness image T'' after the filter processing.
  • the thickness image T'' is a thickness image to which flicker reduction and noise reduction have been applied in a double manner.
  • Spatial filters such as Gaussian filters and median filters, structure-preserving spatial filters such as epsilon filters and Laplacian filters, and temporal filters such as recursive filters can be used for the filtering process of block F1.
  • a block H1 generates an accumulated image A'' using, for example, equation (15) from the high energy image H'' and the low energy image L'' subjected to the blink reduction process, which are output from the block R1.
  • a block MD2 performs material separation processing from the thickness image T'' output from the block F1 and the accumulated image A'' output from the block H1, and performs flickering reduction processing and noise reduction processing to obtain a bone image. B''' is generated. The material separation processing by block MD2 will be described later.
  • a block R3 performs further flickering reduction processing on the bone image B''' obtained by the block MD2 to generate a flickering-reduced bone image B''''. According to the noise reduction processing using the block F1 and the block MD2 as described above, noise accompanying the energy subtraction processing (substance separation) is reduced.
  • Equation (17) By substituting the pixel value A and the thickness T of the accumulated image at a certain pixel into Equation (17) and solving the nonlinear equation, it is possible to obtain the thickness B of the bone at a certain pixel. That is, if the thickness image T′ subjected to noise reduction processing and flickering reduction processing is used as the thickness T, and the accumulated image A′′ output from the block H1 is used as the pixel value A of the accumulated image, Equation (17) is solved. , the bone thickness B representing the bone image B''' is obtained. Since the thickness image has higher continuity than the accumulated image, it does not contain high frequency components. Therefore, even if noise is removed by filtering the thickness image T' using the block F1, the signal component is less likely to be lost.
  • the noise-reduced bone image B''' can be obtained by the block MD2 using the noise-reduced thickness image T'' and the accumulated image A'', which originally has little noise.
  • block MD2 obtains a noise-reduced soft tissue image S''' from the thickness image T'' and the accumulated image A''. It is also possible to
  • the image T'' synthesized with the block MD2 is used, but it is not limited to this.
  • the block MD2 performs noise reduction and blink reduction.
  • a bone image B''' obtained by performing the above may be obtained.
  • the block MD2 is subjected to noise reduction and blink reduction.
  • a soft tissue image S''' may be obtained.
  • the noise reduction processing by the filters of block F2 and block F3 may be omitted.
  • the noise reduction processing by the filter in block F1 may be omitted.
  • the flickering reduction processing used for block R1, block R2, and block R3 may be either a method of multiplying coefficients or a method of adding or subtracting coefficients.
  • blinking can be suppressed most effectively when multiplication correction is performed in blinking reduction processing for the high energy image H and the low energy image L.
  • flickering could be suppressed most effectively when addition and subtraction were performed in flickering reduction processing for the bone image B' and the soft tissue image S'. Therefore, it is preferable to apply correction by multiplication to the blink reduction processing of block R1, and correction by addition and subtraction to the blink reduction processing of blocks R2 and R3.
  • the flickering reduction process is performed in blocks R1 to R3, but the present invention is not limited to this. At least one of block R1, block R2, and block R3 may be subjected to the blink reduction process. However, according to the study of the present inventors, it is preferable from the viewpoint of blinking reduction to perform the blinking reduction processing by the block R1 and the block R2.
  • filter blocks F2 and F3 applied before block MD1 for two-substance separation, and filter block F1 applied after block MD1 It is necessary to optimize the type and strength at the same time. This is because the result of optimizing the two filters independently is not always optimal. For example, if a temporal direction filter or a spatial direction filter is applied twice, the increase rate of the quantum noise of X-rays and the noise accompanying the separation of two substances will not be independent, and the two noise reduction effects may not be integrated.
  • blocks F2 and F3 which are filters applied before block MD1 for two-substance separation apply filters in the temporal direction
  • filters applied after MD1, block F1 apply filters in the spatial direction.
  • a configuration can be used.
  • the filter in the spatial direction or the filter in the temporal direction is doubled, and in that case, it is preferable to make the size of the kernel and the size of the filter coefficients different for both.
  • the filter kernel of block F1 when filtering in the spatial direction twice, it is preferable to configure the filter kernel of block F1 to be larger than the kernels of blocks F2 and F3. This is because the thickness image T has higher spatial continuity than the accumulated image A, the high energy image H, and the low energy image L, for example.
  • the filter coefficient of the block F1 when applying filters in the time direction twice, it is preferable to configure the filter coefficient of the block F1 to be larger than the filter coefficients of the filters of the blocks F2 and F3. This is because, for example, the thickness image T changes less over time than the accumulated image A, the high energy image H, and the low energy image L.
  • block F1 may be configured to apply both temporal and spatial filters
  • blocks F2 and F3 may be configured to simultaneously apply both temporal and spatial filters.
  • a spatial or temporal filter is doubly applied before and after the two-matter separation block MD1
  • a configuration with larger filter coefficients or kernels in the filter block F1 after MD1. can be preferably used.
  • the X-ray imaging device 104 is an indirect radiation sensor using a phosphor, but is not limited to such a form.
  • a direct radiation sensor using a direct conversion material such as CdTe may be used.
  • the passive tube voltage change of the X-ray generator 101 is used (FIG. 3A), or the tube voltage is actively switched (FIG. 3B). is not limited to any form.
  • the energy of the radiation irradiated to the X-ray imaging device 104 may be changed by switching the filter of the X-ray generation device 101 over time.
  • the present invention is not limited to such a form, and the above-described embodiment can be applied even when the X-ray (radiation) has three or more energies.
  • energy subtraction processing using n images corresponding to n energies (n is a natural number of 2 or more), n material separation images are obtained.
  • blink reduction processing By applying the above-described blink reduction processing to n images (moving images) before energy subtraction processing and/or to n separated images (moving images) after energy subtraction processing, separated images with reduced blinking are obtained. .
  • energy subtraction was performed by changing the energy of the radiation irradiated to the X-ray imaging device 104, but the present invention is not limited to such a form.
  • a method of changing the spectrum of radiation detected by the front sensor and the rear sensor may be used.
  • a plurality of images with different energies may be obtained by using a photon counting sensor that counts the number of radiation quanta for each energy.
  • energy subtraction processing was performed using the control computer 103 of the radiation imaging system, but the present invention is not limited to such a form.
  • the control computer 103 may be incorporated into the X-ray imaging device 104 .
  • the image acquired by the control computer 103 may be transferred to another computer to perform energy subtraction processing.
  • a configuration in which an acquired image is transferred to another personal computer (image viewer) via a medical PACS and displayed after energy subtraction processing is preferably used. That is, in each of the above embodiments, it is sufficient to provide radiation images with different energies to the energy subtraction process, and the method for acquiring radiation images with different energies is not limited to the above embodiments.
  • control computer 103 directly acquires an image from the X-ray imaging apparatus 104 and performs energy subtraction processing, but the present invention is not limited to this.
  • a moving image captured by the X-ray imaging apparatus 104 may be stored in an external storage device, and the control computer 103 may read out the moving image from the storage device and perform energy subtraction processing.
  • the present invention supplies a program that implements one or more functions of the above-described embodiments to a system or device via a network or a storage medium, and one or more processors in the computer of the system or device reads and executes the program. It can also be realized by processing to It can also be implemented by a circuit (for example, ASIC) that implements one or more functions.
  • a circuit for example, ASIC
  • 101 X-ray generator
  • 102 X-ray controller
  • 103 Control computer
  • 104 X-ray generator

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Abstract

Ce dispositif de traitement d'image utilise une pluralité d'images animées correspondant à une pluralité de différents types d'énergie pour effectuer un traitement de soustraction d'énergie et générer ainsi des images animées d'images séparées, et corrige la pluralité d'images animées ou les images animées des images séparées de façon à réduire les fluctuations de la valeur de signal entre des trames dans la pluralité d'images animées ou les images animées des images séparées.
PCT/JP2021/047539 2021-02-26 2021-12-22 Dispositif et procédé de traitement d'image, système de radiographie et programme WO2022181022A1 (fr)

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Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JPS60224386A (ja) * 1984-04-23 1985-11-08 Fuji Photo Film Co Ltd サブトラクシヨン画像の濃度補正方法および装置
JP2007050052A (ja) * 2005-08-16 2007-03-01 Canon Inc 放射線撮像装置及びその制御方法
JP2019051322A (ja) * 2018-10-02 2019-04-04 コニカミノルタ株式会社 動態解析システム
WO2020241110A1 (fr) * 2019-05-30 2020-12-03 キヤノン株式会社 Dispositif de traitement des images, procédé de traitement des images et programme
JP2020203083A (ja) * 2019-06-11 2020-12-24 キヤノン株式会社 放射線撮像装置及び放射線撮像システム

Patent Citations (5)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
JPS60224386A (ja) * 1984-04-23 1985-11-08 Fuji Photo Film Co Ltd サブトラクシヨン画像の濃度補正方法および装置
JP2007050052A (ja) * 2005-08-16 2007-03-01 Canon Inc 放射線撮像装置及びその制御方法
JP2019051322A (ja) * 2018-10-02 2019-04-04 コニカミノルタ株式会社 動態解析システム
WO2020241110A1 (fr) * 2019-05-30 2020-12-03 キヤノン株式会社 Dispositif de traitement des images, procédé de traitement des images et programme
JP2020203083A (ja) * 2019-06-11 2020-12-24 キヤノン株式会社 放射線撮像装置及び放射線撮像システム

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