WO2018123403A1 - Système d'imagerie par rayonnement - Google Patents

Système d'imagerie par rayonnement Download PDF

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WO2018123403A1
WO2018123403A1 PCT/JP2017/042741 JP2017042741W WO2018123403A1 WO 2018123403 A1 WO2018123403 A1 WO 2018123403A1 JP 2017042741 W JP2017042741 W JP 2017042741W WO 2018123403 A1 WO2018123403 A1 WO 2018123403A1
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signal
signal value
imaging system
radiation imaging
value
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PCT/JP2017/042741
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English (en)
Japanese (ja)
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晃介 照井
野田　剛司
佳士 町田
貴司 岩下
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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 or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment

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  • the present invention relates to a radiation imaging system.
  • a radiation imaging apparatus using a flat panel detector (FPD) formed of a semiconductor material is known as an imaging apparatus used for medical image diagnosis and nondestructive inspection using radiation (X-rays).
  • the radiation imaging apparatus can be used as a digital imaging apparatus for still images and moving images, for example, in medical image diagnosis.
  • the FPD for example, there are an integration type sensor and a photon counting type sensor.
  • the integral type sensor measures the total amount of charge generated by the incidence of radiation.
  • the photon counting type sensor discriminates the energy (wavelength) of incident radiation and counts the number of times of detection of radiation for each of a plurality of energy levels. That is, since the photon counting type sensor has energy resolution, the diagnostic ability can be improved as compared with the integration type sensor. However, since the number of incident radiation quanta is enormous, a high operating speed is required to count them individually. For this reason, it has been difficult to realize a photon counting type sensor with a large area FPD.
  • Patent Document 1 proposes a radiation imaging apparatus having energy resolution by estimating the number of radiation quanta and the average value of energy using average image density information and image density dispersion information for each predetermined region. Has been. With the method of Patent Document 1, a sensor having energy resolution can be realized even at a lower operating speed than a photon counting type sensor.
  • the present invention has been made with the above problem recognition as an opportunity, and an object thereof is to provide an advantageous technique for more accurately obtaining the energy of radiation quanta.
  • One aspect of the present invention relates to a radiation imaging system, and the radiation imaging system includes a scintillator that converts radiation quanta into light and a plurality of photoelectric conversion elements, and a plurality of signals provided from the detector.
  • an advantageous technique is provided for more accurately obtaining the energy of radiation quantum.
  • the figure which shows the structure of the radiation imaging system of one embodiment of this invention The figure which shows the structural example of a detector. The figure which shows the structural example of a pixel. The figure which shows the operation example of 1 frame cycle. The figure explaining the principle which calculates
  • the figure which illustrates the process performed by the process part of a control apparatus The figure which shows the pixel value of one pixel in the some digital image signal output from a radiation imaging device in case there exists blinker noise, and distribution of this pixel value, respectively.
  • FIG. 1 shows the configuration of a radiation imaging system according to one embodiment of the present invention.
  • the radiation imaging system 1 captures an image formed by radiation emitted from the radiation source 11 and transmitted through the subject.
  • the radiation may typically be X-rays, but may be alpha rays, beta rays, gamma rays, and the like.
  • the radiation imaging system 1 can include a detector 101 and a control device 13.
  • the radiation imaging system 1 includes a radiation imaging apparatus 10, a control apparatus 13, an exposure control apparatus 12, and a radiation source (for example, an X-ray source) 11.
  • the exposure control device 12 may be incorporated in the radiation source 11 or the control device 13.
  • the radiation imaging apparatus 10 can include a detector 101, an output unit 105, a power supply unit 106, and an imaging control unit 107.
  • the detector 101 includes a scintillator 210 that converts radiation quanta into light and a plurality of photoelectric conversion elements 201.
  • the detector 101 includes a pixel array 102 in which a plurality of pixels 20 are arranged so as to form a plurality of rows and a plurality of columns, and one pixel 20 is configured by one photoelectric conversion element 201, and a plurality of pixels 20 can share the scintillator 210.
  • the photoelectric conversion element 201 converts light (visible light) converted from radiation quanta by the scintillator 210 into electric charge.
  • the photoelectric conversion element 201 converts radiation quanta transmitted through the scintillator 210 and incident on the photoelectric conversion element 201 into electric charges.
  • the radiation quanta are converted into electric charges in the photoelectric conversion element 201 that passes through the scintillator 210 and enters the photoelectric conversion element 201, blinker noise as described above may occur.
  • the radiation imaging apparatus 10 can include a drive circuit 103 that drives the pixel array 102 and a readout circuit 104 that reads a signal from the driven pixel array 102.
  • the control device 13 gives control signals to the radiation imaging device 10 and the exposure control device 12 based on imaging information input from a photographer (not shown) via a control console (not shown) of the control device 13.
  • the exposure control device 12 can control the operation of emitting radiation from the radiation source 11 and the operation of an irradiation field stop mechanism (not shown).
  • the imaging control unit 107 of the radiation imaging apparatus 10 receives each control signal from the control apparatus 13 and controls each part of the radiation imaging apparatus 10. An image of radiation emitted from the radiation source 11 controlled by the exposure control device 12 and transmitted through the subject is formed on the detector 101, and the detector 101 outputs a signal corresponding to the image.
  • the output unit 105 performs image processing such as offset correction after AD conversion of the signal output from the detector 101, and transmits the processed signal to the control device 13.
  • image processing such as offset correction after AD conversion of the signal output from the detector 101
  • known wireless communication or wired communication can be applied to the transmission.
  • the transmitted signal is processed by the processing unit 131 of the control device 13 and can be displayed on a display unit (not shown) of the control device 13.
  • FIG. 3 shows a configuration example of each pixel 20.
  • the pixel 20 can include a pixel circuit unit 202 for reading a signal from the photoelectric conversion element 201.
  • the photoelectric conversion element 201 can typically be a photodiode.
  • the pixel circuit unit 202 includes an amplifier circuit unit 204, a clamp circuit unit 206, a sample hold circuit unit 207, and a selection circuit unit 208.
  • the photoelectric conversion element 201 includes a charge storage unit, and the charge storage unit is connected to the gate of the MOS transistor 204a of the amplifier circuit unit 204.
  • the source of the MOS transistor 204a is connected to the current source 204c through the MOS transistor 204b.
  • the MOS transistor 204a and the current source 204c constitute a source follower circuit.
  • the MOS transistor 204b is an enable switch that is turned on when the enable signal EN supplied to the gate thereof is set to the active level by the drive circuit 103 to bring the source follower circuit into an operating state.
  • the charge storage unit of the photoelectric conversion element 201 and the gate of the MOS transistor 204a constitute a common node, which converts the charge stored in the charge storage unit into a voltage.
  • Functions as a charge-voltage converter. That is, the voltage V ( Q / C) determined by the charge Q stored in the charge storage unit and the capacitance value C of the charge voltage conversion unit appears in the charge voltage conversion unit.
  • the charge-voltage converter is connected to the reset potential Vres via the reset switch 203.
  • the reset signal PRES is set to the active level by the drive circuit 103, 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 unit 206 clamps the noise output from the amplifier circuit unit 204 by the clamp capacitor 206a according to the reset potential of the charge-voltage conversion unit. That is, the clamp circuit unit 206 is a circuit for canceling this noise from the signal output from the source follower circuit in accordance with the electric charge generated by the 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 an inactive level and turning off the MOS transistor 206b after the drive circuit 103 sets the clamp signal PCL to an active level to turn on the MOS transistor 206b. The output side of the clamp capacitor 206a is connected to the gate of the MOS transistor 206c.
  • the source of the MOS transistor 206c is connected to the current source 206e via the MOS transistor 206d.
  • the MOS transistor 206c and the current source 206e constitute a source follower circuit.
  • the MOS transistor 206d is an enable switch that is turned on when the enable signal EN0 supplied to the gate thereof is set to the active level by the drive circuit 103 to bring the source follower circuit into an operating state.
  • a signal output from the clamp circuit unit 206 according to the electric charge generated by the photoelectric conversion in the photoelectric conversion element 201 is written as an optical signal into the capacitor 207Sb via the switch 207Sa when the optical signal sampling signal TS becomes an active level. It is.
  • a signal output from the clamp circuit unit 206 when the MOS transistor 206b is turned on immediately after resetting the potential of the charge voltage conversion unit is noise. This noise is written into the capacitor 207Nb via the switch 207Na when the noise sampling signal TN becomes an active level. This noise includes an offset component of the clamp circuit unit 206.
  • the switch 207Sa and the capacitor 207Sb constitute a signal sample / hold circuit 207S
  • the switch 207Na and the capacitor 207Nb constitute a noise sample / hold circuit 207N
  • the sample hold circuit unit 207 includes a signal sample hold circuit 207S and a noise sample hold circuit 207N.
  • a signal (radiation signal) held in the capacitor 207Sb is output to the signal line 25S via the MOS transistor 208Sa and the row selection switch 208Sb.
  • a signal (noise) held in the capacitor 207Nb is output to the signal line 25N via the MOS transistor 208Na and the row selection switch 208Nb.
  • the MOS transistor 208Sa forms a source follower circuit with a constant current source (not shown) provided on the signal line 25S.
  • the MOS transistor 208Na constitutes a constant current source (not shown) provided on the signal line 25N and a source follower circuit.
  • the MOS transistor 208Sa and the row selection switch 208Sb constitute a signal selection circuit unit 208S
  • the MOS transistor 208Na and the row selection switch 208Nb constitute a noise selection circuit unit 208N.
  • the selection circuit unit 208 includes a signal selection circuit unit 208S and a noise selection circuit unit 208N.
  • the readout circuit 104 generates a radiation signal from which noise components have been removed by amplifying the difference between the radiation signal and noise output to the signal lines 25S and 25N, respectively.
  • the pixel 20 may include an addition switch 209 ⁇ / b> S that adds the optical signals of the plurality of adjacent pixels 20.
  • the addition mode signal ADD becomes an active level, and the addition switch 209S is turned on.
  • the capacitors 207Sb of the adjacent pixels 20 are connected to each other by the addition switch 209S, and the optical signals are averaged.
  • the pixel 20 may include an addition switch 209N that adds noises of a plurality of adjacent pixels 20. When the addition switch 209N is turned on, the capacitors 207Nb of the adjacent pixels 20 are connected to each other by the addition switch 209N, and noise is averaged.
  • the adding unit 209 includes an addition switch 209S and an addition switch 209N.
  • the pixel 20 may have a sensitivity changing unit 205 for changing the sensitivity.
  • the pixel 20 may include, for example, a first sensitivity change switch 205a and a second sensitivity change switch 205'a, and circuit elements associated therewith.
  • 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 an 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 set to an active level, and the MOS transistor 204'a may be operated as a source follower in addition to the MOS transistor 204a.
  • FIG. 3 illustrates a procedure for obtaining one frame (one sheet) of radiation image in the radiation imaging system 1.
  • the horizontal axis represents time.
  • One frame cycle (one frame period) for obtaining a radiographic image of one frame can include an irradiation period XW, a readout period XR, a non-irradiation period FW, and a readout period FR.
  • the irradiation period XW is a period during which the detector 101 is irradiated with radiation.
  • the readout period XR is a period in which the readout circuit 104 reads out a radiation signal from the detector 101.
  • the length of the non-irradiation period FW can be set to be the same as the length of the irradiation period XW.
  • the non-irradiation period FW is a period in which charges are accumulated in each pixel 20 (photoelectric conversion element 201) in a state where radiation is not irradiated in order to generate offset information.
  • the readout period FR is a period in which the readout circuit 104 reads offset information from the detector 101.
  • the radiation imaging system 1 can obtain a plurality of frames of radiation image signals by repeating one frame cycle shown in FIG. 4 a plurality of times.
  • the signal read from the detector 101 by the reading circuit 104 is provided to the processing unit 131 of the control device 13 via the output unit 105.
  • the processing unit 131 of the control device 13 processes the radiation image signal provided from the radiation imaging apparatus 10. Specifically, the processing unit 131 removes the offset by calculating the difference between the radiation signal read from the detector 101 in the readout period XR and the signal read from the detector 101 in the readout period FR. Obtained radiographic image signals are obtained.
  • the processing unit 131 is, for example, a PLD (abbreviation of Programmable Logic Device) such as an FPGA (abbreviation of Field Programmable Gate Array), or an ASIC (Abbreviation of Application Specific Integrated, or an abbreviation of a generalized circuit program). It can be constituted by a computer or a combination of all or part of them. Or the process part 131 may be comprised as a part of function comprised by the program provided to the computer or CPU which comprises the control apparatus 13. FIG.
  • the processing unit 131 obtains the energy of radiation quanta (for example, X-ray photons) incident on the detector 101 based on a signal provided from the detector 101.
  • the processing unit 131 further obtains an energy-resolved radiation image.
  • the energy-resolved radiographic image is a radiographic image having the magnitude of energy as a gradation value.
  • FIG. 5 shows the principle for obtaining the energy of X-ray photons as radiation quanta.
  • X-ray photons that are an example of radiation quanta are absorbed by the scintillator
  • visible photons are generated by the scintillator.
  • the number of visible photons generated depends on the energy of X-ray photons absorbed by the scintillator. Specifically, the greater the energy of X-ray photons, the more visible light photons are generated in the scintillator.
  • Visible light photons are absorbed by the photoelectric conversion element, and charges are generated in the photoelectric conversion element accordingly.
  • the number of generated charges (charge amount) depends on the number of visible photons absorbed by the photoelectric conversion element.
  • the output circuit can convert an analog signal corresponding to the number of charges generated by the photoelectric conversion element into a digital signal and output the digital signal.
  • the value of a digital signal output from the output circuit in response to an X-ray photon having a certain energy is 30 LSB
  • the value of a digital signal output in response to an X-ray photon having a higher energy is 100 LSB. is there. Therefore, if a digital signal corresponding to the amount of charge generated in the photoelectric conversion element is acquired every time one X-ray photon is absorbed by the scintillator, the energy of the X-ray photon can be identified from the value. is there.
  • LSB is used as a quantization unit for analog / digital conversion.
  • 30LSB means 30 quantization units. In this way, an energy-resolved radiation image can be obtained by discriminating the energy of incident radiation quanta.
  • FIG. 6 is a conceptual diagram for explaining the principle of estimating the average value of radiation quantum energy.
  • a plurality of digital image signals are acquired by irradiating the detector with radiation for a predetermined period of time through the subject a plurality of times.
  • the radiation irradiated for a predetermined period is constant and the subject does not move.
  • FIG. 6 shows a time series of digital signal values (hereinafter referred to as pixel values or signal values) obtained by selecting any one pixel from the obtained digital image signals and obtained from the selected pixels. Yes.
  • the pixel value should ideally be constant.
  • the pixel value actually varies in time series. This variation includes quantum noise.
  • Quantum noise is generated when the number of radiation quanta per unit time (for example, the number of X-ray photons) varies.
  • the variation in the number of radiation quanta is a Poisson distribution, which is a discrete probability distribution with a specific random variable that counts discrete events that occur at a given time interval, when considered as the probability of occurrence per unit time for discrete events.
  • the number of radiation quanta actually irradiated to any one pixel per unit time is 12 Five, 13, 11, etc. vary.
  • the value of the digital signal output according to one radiation quantum of a certain energy is 30 LSB
  • the actual pixel value is 360 LSB, 150 LSB, 390 LSB, It varies such as 330LSB.
  • the expected value of the pixel value is 300 LSB
  • the variation hereinafter referred to as dispersion
  • the expected value of the quantum number of radiation quanta irradiated to an arbitrary pixel per unit time is 3, and the value of a digital signal output according to one radiation quantum of a certain energy is 100 LSB.
  • the expected value of the pixel value is 300 LSB, and the variation (hereinafter referred to as dispersion) is 30000 LSB.
  • the dispersion of the pixel values is larger in the pixels formed by the radiation quanta having a large energy.
  • the energy of radiation quanta such as X-ray photons can be estimated.
  • the radiation imaging apparatus 10 (detector 101) of the radiation imaging system 1 is irradiated with radiation T (T is a natural number of 2 or more) times, and T digital image signals are acquired from the radiation imaging apparatus 10.
  • T the pixel value of a certain pixel of the t-th (t is a natural number of 2 or more and T or less) digital image signal
  • I the total quantum number of the radiation quanta absorbed by reaching the pixel
  • n Var the sample dispersion of the quantum number of the radiation quantum that reaches the pixel of the digital image signal and is absorbed
  • n Var ⁇ [ ⁇ I (t) / E ⁇ n Ave ⁇ 2 ] / T (3)
  • the expected value and the variance are equal to the parameter ⁇ .
  • the arithmetic mean approaches the expected value
  • the sample variance approaches the variance. Therefore, when the number of samples is sufficiently large (preferably infinite) and the arithmetic mean n Ave of the quantum numbers of the radiation quanta and the sample variance of the quantum numbers of the radiation quanta are approximated to be equal to n Var , the equation (2) And the following expression (4) is derived from the assumption that the expression (3) is equal.
  • the sample variance I Var of the pixel value is expressed by the following formula (6) from the sample variance of the quantum number of the radiation quantum n Var .
  • FIG. 7A illustrates the pixel value I (t) of one pixel in a plurality of digital image signals output from the radiation imaging apparatus 10 when there is no blinker noise.
  • FIG. 7B shows the distribution (histogram) of the pixel values I (t) in FIG. 7A. This distribution follows the Poisson distribution.
  • FIG. 8A illustrates the pixel value I (t) of one pixel in a plurality of digital image signals output from the radiation imaging apparatus 10 when there is blinker noise.
  • FIG. 8B shows the distribution (histogram) of the pixel values I (t) in FIG. 8A. This distribution follows the Poisson distribution.
  • blinker noise occurs, the pixel value becomes larger than the normal pixel value. This is because when blinker noise occurs, a component due to radiation quanta incident on the photoelectric conversion element 201 (component due to blinker noise) is added to a component due to visible light incident on the photoelectric conversion element 201.
  • the distribution of the pixel value I (t) can be a distribution deviating from the Poisson distribution as illustrated in FIG. 8B. Even if the arithmetic mean obtained from the pixel value I (t) having a distribution deviating from the Poisson distribution due to blinker noise is obtained based on I Ave and the sample variance I Var and the energy E of the radiation quantum is obtained according to the equation (7), Energy E cannot be obtained. Therefore, in the present embodiment, the processing unit 131 reduces the influence of blinker noise in a signal value group including a plurality of signal values (pixel values) provided from the radiation imaging apparatus 10 (detector 101). The signal value group may be modified to generate a modified signal value group.
  • the processing unit 131 can be configured to obtain the energy of radiation quanta incident on the detector 101 based on the correction signal group.
  • the processing unit 131 includes a feature amount indicating a feature of the distribution of the signal values constituting the modified signal value group (for example, arithmetic mean I AVE and sample variance of the signal values (pixel values) constituting the modified signal value group). It may be configured to determine the energy E based on I Var ).
  • the processing unit 131 excludes a signal selected according to a predetermined criterion from a plurality of signal values provided from the radiation imaging apparatus 10 (detector 101), or uses the signal value as an original signal value.
  • a modified signal value group is obtained by changing to a smaller signal value. More specifically, for example, the processing unit 131 excludes a signal value exceeding the threshold Tth from a plurality of signal values provided from the radiation imaging apparatus 10 (detector 101), or uses the signal value as an original signal value.
  • a modified signal value group is obtained by changing to a smaller signal value.
  • the signal value smaller than the original signal value is, for example, a representative value (for example, average value, median value, mode value) of a plurality of signals constituting the signal value group, or a signal corresponding to the threshold value Tth.
  • the threshold value Tth can be determined based on experiments or the like. Alternatively, the threshold value Tth is obtained by obtaining a corrected signal value group using the tentatively determined threshold value Tth, and evaluating the divergence amount between the distribution of the plurality of signal values constituting the corrected signal value group and the Poisson distribution. May be determined by optimizing the threshold Tth until the value falls within the reference value.
  • FIG. 9 illustrates processing executed by the processing unit 131.
  • the processing unit 131 calculates the difference between the radiation signal read from the detector 101 in the readout period XR and the signal read from the detector 101 in the readout period FR, thereby removing the offset.
  • a radiographic image signal (digital image signal) is acquired.
  • the processing unit 131 may be configured to acquire a plurality (a plurality of frames) of radiographic image signals.
  • the plurality of radiographic image signals include, for example, a signal group including a plurality of signals for each pixel.
  • step S902 the processing unit 131 generates a correction signal value group for each pixel by processing each signal value group in the plurality of radiation image signals acquired in step S901 as described above.
  • step S903 the processing unit 131 obtains a feature amount from the correction signal value group for each pixel generated in step S902.
  • the feature amount is an amount that indicates the characteristics of the distribution of a plurality of signal values constituting the modified signal value group. For example, the arithmetic mean I AVE and the sample variance I of the signal values (pixel values) constituting the modified signal value group. Var ).
  • step S904 the processing unit 131 can obtain energy E according to Expression (7) based on the feature amount obtained in step S903. By obtaining the energy E for each pixel, an energy-resolved radiation image can be obtained.
  • the method for obtaining the corrected signal value group various methods can be adopted in addition to the method using the threshold as described above. For example, there is a method of sorting a plurality of signal values constituting a signal value group in descending order and removing a higher-order signal value in the sorting result.
  • the number of signal values to be removed can be determined based on the occurrence frequency of blinker noise. Since blinker noise is generated when the radiation that has passed through the scintillator 210 is absorbed by the photoelectric conversion element 201, the blinker noise is based on the amount of radiation incident on the scintillator 210, the thickness of the scintillator 210, the quantum efficiency of the photoelectric conversion element 201, and the like. The frequency of occurrence can be estimated.
  • the scintillator 210 is made of CsI having a thickness of 1000 ⁇ m, made of a photoelectric conversion element 201Si, and the radiation is X.
  • the radiation is X.
  • about 20% of the X-rays incident on the scintillator 210 are transmitted, and less than 1% is absorbed by the photoelectric conversion element (having little sensitivity to light with a wavelength of pm order), and blinker noise.
  • the photoelectric conversion element having little sensitivity to light with a wavelength of pm order
  • the correction signal value group is a signal value that exceeds the threshold value among the signal values provided from the radiation imaging apparatus 10. May be obtained by subtracting the blinker noise component from The blinker noise component can be determined by subtracting the signal value when no blinking occurs as shown in FIG. 7A from the signal value when the blinker noise occurs as shown in FIG. 8A.
  • the second embodiment is a correction that is effective when the blinker noise component is small with respect to the quantum variation, and the signal value when the blinker noise is generated is close to the signal value when the blinker noise is not generated.
  • a method for generating a signal value group (step S902) is provided.
  • the signal value including the blinker noise component has a signal value larger than a representative value (median value, mode value, average value, etc.) of a plurality of signal value distributions constituting the signal value group.
  • a modified signal value group is generated from the signal value group by using.
  • FIG. 11 visually illustrates a method of generating a correction signal value group (step S902) in the second embodiment.
  • step S902 a distribution of a plurality of signal values constituting a signal value group provided from the radiation imaging apparatus 10 is illustrated.
  • the right side of FIG. 11 illustrates a distribution of a plurality of signal values constituting the modified signal value group generated by the processing unit 131 in step S902 based on the left signal value group.
  • the signal value including the blinker noise component is larger than the representative value of the plurality of signal values constituting the signal value group according to the experiment by the present inventor. Therefore, a blinker noise component is included in the signal value in the section between the median value and the minimum value (first section between the representative value and the minimum value) in the distribution of the plurality of signal values constituting the signal value group. Most likely not. Therefore, the processing unit 131 determines the representative value and the plurality of signals based on the signal value in the first section between the representative value of the plurality of signal values constituting the signal value group and the minimum value of the plurality of signal values. The signal value in the modified signal value group corresponding to the second interval between the maximum values in the values may be determined. For example, the processing unit 131 can generate the corrected signal value group so that the distribution of the plurality of signal values constituting the corrected signal value group is symmetric with the representative value as the symmetry axis.
  • step S902 an operation example of the processing unit 131 for realizing the correction signal value group generation method (step S902) in the second embodiment will be described with reference to FIGS. 12A and 12B.
  • the median value is adopted here as the representative value, there is no significant difference even if the mode value or the average value is adopted as the representative value.
  • the processing unit 131 sorts a plurality of signal values constituting the signal value group in ascending order, and generates a data string including data [0]... Data [n].
  • the processing unit 131 calculates a difference ⁇ 0 to ⁇ m ⁇ 1 between the median value Xm and a signal value smaller than the median value Xm.
  • the processing unit 131 determines a signal value larger than the median value Xm among a plurality of signal values constituting the signal value group based on the differences ⁇ 0 to ⁇ m ⁇ 1 .
  • a corrected signal value group is generated.
  • the processing unit 131 corrects the signal value distribution in the second interval between the median value and the maximum value so that the signal value distribution in the first interval is smaller than the median value. Is generated. More specifically, the processing unit 131 generates a correction signal group by adding the median Xm and ⁇ m ⁇ i to data [m + i] that is the signal value of the second section.
  • the third embodiment is a modified signal value group that is effective when the blinker noise component is small with respect to quantum variation and the signal value when the blinker noise is generated is close to the signal value when the blinker noise is not generated.
  • a generation method (step S902) is provided.
  • the third embodiment can also be understood as a modification of the second embodiment.
  • FIG. 13 visually illustrates a method of generating a correction signal value group (step S902) in the third embodiment.
  • reference numeral 1301 denotes a distribution of a plurality of signal values constituting a signal value group provided from the radiation imaging apparatus 10
  • reference numeral 1302 denotes a candidate signal value generated from 1301 by the same method as in the second embodiment.
  • Reference numeral 1303 denotes a modified signal value group generated based on the signal value group 1301 and the candidate signal value group 1302 generated based on the signal value group 1301.
  • a modified signal value group similar to the candidate signal value group is generated.
  • the processing unit 131 uses the signal value group 1301 for a section where the difference between the signal value constituting the signal value group 1301 and the signal value constituting the candidate signal value group 1302 is smaller than a predetermined value.
  • the signal values of the modified signal value group 1303 are used as they are without changing the constituent signal values.
  • the processing unit 131 determines the signal value constituting the signal value group 1301 as The candidate signal value group 1302 is replaced with a candidate signal value.
  • the energy of the radiation quantum can be obtained more faithfully by the actual signal value (the signal value detected by the detector 101). it can.
  • the occurrence frequency of blinker noise can be reduced by incorporating an FOP (fiber optic plate) between the scintillator 210 and the photoelectric conversion element 201.
  • the FOP is a glass plate in which fibers with a pitch of several um are gathered, and absorbs radiation while transmitting visible light without being scattered.
  • a technique for reducing the influence of blinker noise as described above is useful even when FOP is incorporated.

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Abstract

Cette invention concerne un système d'imagerie par rayonnement comprenant : un détecteur qui comprend une pluralité d'éléments de conversion photoélectriques et un scintillateur pour convertir des quanta de rayonnement en lumière ; et une unité de traitement qui détermine l'énergie des quanta de rayonnement incidents sur le détecteur, en fonction d'un groupe de valeurs de signal modifié obtenues par modification d'un groupe de valeurs de signal constitué d'une pluralité de valeurs de signal fournies par le détecteur de façon à réduire l'influence du bruit de clignotant sur le groupe de valeurs de signal.
PCT/JP2017/042741 2016-12-27 2017-11-29 Système d'imagerie par rayonnement WO2018123403A1 (fr)

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JPH06125887A (ja) * 1992-10-15 1994-05-10 Hamamatsu Photonics Kk 医療用ｘ線画像処理装置
JP2009285356A (ja) * 2008-05-30 2009-12-10 Institute Of National Colleges Of Technology Japan 医療用撮影システム、画像処理装置、画像処理方法、およびプログラム
JP2010263961A (ja) * 2009-05-12 2010-11-25 Canon Inc Ｘ線画像撮影装置およびｘ線画像撮影装置の制御方法
JP2014128456A (ja) * 2012-12-28 2014-07-10 Toshiba Corp Ｘ線ｃｔ装置及び制御プログラム

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JPH06125887A (ja) * 1992-10-15 1994-05-10 Hamamatsu Photonics Kk 医療用ｘ線画像処理装置
JP2009285356A (ja) * 2008-05-30 2009-12-10 Institute Of National Colleges Of Technology Japan 医療用撮影システム、画像処理装置、画像処理方法、およびプログラム
JP2010263961A (ja) * 2009-05-12 2010-11-25 Canon Inc Ｘ線画像撮影装置およびｘ線画像撮影装置の制御方法
JP2014128456A (ja) * 2012-12-28 2014-07-10 Toshiba Corp Ｘ線ｃｔ装置及び制御プログラム

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