JP6470986B2 - Radiation detector and radiation tomography apparatus - Google Patents

Description

  The present invention relates to a technique for improving a radiation detector used for radiography.

  Conventionally, an X-ray detector used in an X-ray computed tomography (CT) apparatus accumulates a large amount of X-ray detection signals per view and uses the total current amount as an output signal. Such an output structure is called a current mode.

  On the other hand, an X-ray detector using a photon counting method receives a single X-ray irradiated for each view and uses a pulsed current generated by the incident X-ray as an output signal. Such an output structure is called a pulse mode.

  Generally, current mode detectors and pulse mode detectors have different structures. The former has a structure in which X-rays are converted into light by a scintillator and then the amount of light is detected by a photodiode (photo-diode). This structure is called an indirect conversion method. The latter has a structure in which X-rays are directly captured by a semiconductor detector, and this structure is called a direct conversion system (see, for example, Patent Document 1, Abstract, etc.).

JP 2014-210047 A

  In the pulse mode, the X-ray energy spectrum can be used, that is, because spectral imaging can be performed, an image can be generated with a richer amount of information than in the current mode. There is. On the other hand, in the pulse mode, a direct X-ray conversion method using a semiconductor detector is generally employed. In this case, the output of the detector is not as stable as the current mode, the count rate characteristics are poor (when the X-ray dose is too high, the output overlaps in the time axis direction, so-called pile-up occurs), many calibrations before and after imaging Many issues remain, such as the need for calibration. Therefore, the current mode is still the mainstay.

  In view of the above circumstances, there is a demand for a technique that can enjoy the merit of the pulse mode as needed while maintaining the image quality and workflow in the current mode with a proven record.

The invention of the first aspect
A radiation detector in which a plurality of radiation detection elements are arranged in at least one direction,
Each of the plurality of radiation detection elements is
A light conversion element that converts incident radiation into photons;
A first photoelectric conversion element that is provided on a first light emission surface of the light conversion element, converts the incident photons into an electrical accumulation signal, and outputs the electrical accumulation signal;
A second photoelectric conversion element that is provided on a second light emission surface different from the first light emission surface of the light conversion element and converts the incident photons into an electrical independent signal and outputs the signal. And a radiation detector comprising:

The invention of the second aspect is
The light conversion element includes a scintillator,
The first photoelectric conversion element includes a photodiode;
The radiation detector according to the first aspect described above, wherein the second photoelectric conversion element includes Si-PM (Silicon Photomultiplier).

The invention of the third aspect is
The first light emitting surface is a bottom surface;
The radiation detector according to the first aspect or the second aspect, wherein the second light emission surface is a side surface.

  Here, the “bottom surface side” means the bottom surface side when the radiation incident direction is the vertical direction. Further, the “side surface side” means the side surface side when the radiation incident direction is the vertical direction.

The invention of the fourth aspect is
A radiation imaging apparatus including the radiation detector according to any one of the first to third aspects is provided.

In a fifth aspect of the invention, the intensity of the radiation detection signal in the radiation detection element is determined based on the magnitude of the accumulated signal output from the first photoelectric conversion element included in the radiation detection element. A first determining means;
A second determination that determines the intensity of a radiation detection signal in the radiation detection element based on a count value within a predetermined time of the independent signal output from the second photoelectric conversion element included in the radiation detection element. And a radiation imaging apparatus according to the fourth aspect of the present invention.

The invention of the sixth aspect is
First reconstruction means for reconstructing a first image based on first projection data based on the intensity of the radiation detection signal determined by the first determination means;
The second reconstructing means for reconstructing the second image based on the second projection data based on the intensity of the radiation detection signal determined by the second determining means. A radiation imaging apparatus is provided.

The invention of the seventh aspect
The second determination means determines the intensity of a detection signal of radiation having specific energy in the radiation detection element based on a count value within a certain time of a signal having a specific wave height among the independent signals,
The radiation imaging apparatus according to the sixth aspect, wherein the second reconstruction means reconstructs a second image based on projection data based on the intensity of a detection signal of radiation having the specific energy.

The invention of the eighth aspect
The radiation imaging apparatus according to the sixth aspect or the seventh aspect, further comprising display means for displaying the first image and the second image in an overlapping manner.

The invention of the ninth aspect is
Control means for controlling the radiation source so as to irradiate and image radiation with a first dose and radiation with a second dose lower than the first dose;
The first determining means determines the intensity of the detection signal of the radiation based on the accumulated signal at the time of irradiation of the radiation with the first dose;
The second determination means determines the intensity of the detection signal of the radiation based on the independent signal at the time of irradiation with the radiation with the second dose, any one of the fifth to eighth aspects A radiographic apparatus according to the aspect is provided.

The invention of the tenth aspect is
The imaging is radiation tomography,
The radiation tomography apparatus according to the ninth aspect, wherein the control unit controls the radiation dose to be irradiated to switch between the first dose and the second dose for each view.

  According to the invention of the above aspect, the merits of the pulse mode can be enjoyed as necessary while maintaining the image quality and workflow in the current mode with a proven record.

1 is a diagram schematically showing a configuration of an X-ray CT apparatus according to an embodiment of the invention. It is a figure which shows the structure of the X-ray detector and collimator apparatus (collimator) which concern on this embodiment. It is an expanded sectional view which shows the structure of a X-ray detection element. It is a figure which shows an example of the time chart (time chart) of the tube current at the time of imaging | photography. It is a flowchart (flowchart) which shows an example of the flow of operation | movement in the X-ray CT apparatus which concerns on this embodiment.

  Embodiments of the invention will be described below.

(First embodiment)
FIG. 1 is a diagram schematically showing the configuration of an X-ray CT apparatus according to an embodiment of the invention.

  The X-ray CT apparatus 1 includes an imaging table 2, a scanning gantry 3, and an operation console 4.

  The imaging table 2 conveys the placed subject 91 to the cavity 3 </ b> B of the scanning gantry 3. The cavity 3B becomes an imaging space.

  The scanning gantry 3 includes an X-ray tube 31, an X-ray detector 32, a collimator device 33, and a data acquisition unit (data acquisition unit) 34.

  The X-ray tube 31 and the X-ray detector 32 are arranged to face each other with the cavity 3B interposed therebetween. The X-ray tube emits X-rays 81 toward the subject 91 arranged in the cavity 3B. The X-ray detector 32 detects the transmitted X-ray of the subject 91 and outputs a signal.

  The collimator device 33 is disposed between the cavity 3 </ b> B and the X-ray detector 32. The collimator device 33 removes scattered radiation incident on the X-ray detector 32.

  The X-ray tube 31, the X-ray detector 32, and the collimator device 33 are rotatably provided around the cavity 3B.

  The data collection unit 34 collects projection data of the subject 91 based on an output signal from the X-ray detector 32 obtained when the subject 91 is imaged. The data collection unit 34 is also generally called DAS (Data Acquisition System).

  The operation console 4 receives various operations from the operator 92. The operation console 4 processes various data and controls each unit. The operation console 4 includes a shooting mode selection unit 41, a shooting condition setting unit 43, a shooting control unit 44, an image reconstruction unit 45, and a display control unit 46. Note that these functional units of the operation console 4 can be realized, for example, by causing an arithmetic processing unit of a computer to execute a predetermined program.

  The shooting mode selection unit 41 selects a shooting mode based on the operation of the operator 92. The imaging mode (imaging mode) is provided with a plurality of modes depending on the detection method of the transmitted X-ray of the subject 91. In this example, three modes are prepared: a current mode based on the conventional method, a pulse mode based on the photon counting method, and a current & pulse mode based on both. The operator 92 performs an operation of designating any desired mode from these. The shooting mode selection unit 41 selects a shooting mode based on the operation.

  The imaging condition setting unit 43 sets the tube voltage and tube current of the X-ray tube 31, the imaging range of the subject 91, the gantry rotation speed, the slice thickness, the reconstruction function, and the like based on the operation of the operator 92. I do.

  The imaging control unit 44 controls each unit to perform imaging of the subject 91 according to the set imaging conditions.

  The image reconstruction unit 45 reconstructs an image based on the projection data collected by the data collection unit 34. For the image reconstruction, for example, a filtered back-projection method, a three-dimensional image reconstruction method, or the like can be used as appropriate.

  The display control unit 46 controls a monitor (not shown) so as to display various images and character information on the screen.

  Here, the X-ray detector 32 and the collimator device 33 will be described in detail.

  FIG. 2 is a diagram showing a configuration of the X-ray detector 32 and the collimator device 33 according to the present embodiment.

  As shown in FIG. 2, the X-ray detector 32 includes a plurality of X-ray detection elements 32 i arranged in a two-dimensional matrix in a channel direction CH and a slice direction SL. Have. The channel direction is the spreading direction of fan beam X-rays irradiated to the subject 91, and the slice direction is the thickness direction of the fan beam X-rays or the body axis direction of the subject 91. . The X-ray detection element 32i has a width Δa in the channel direction CH and a width Δb in the slice direction SL. The width Δa and the width Δb are, for example, 1 mm. The X-ray incident surface of the X-ray detector 32 is formed by the X-ray incident surface of the X-ray detection element 32i. For example, about 1000 X-ray detection elements 32i are arranged in the channel direction CH and 64 in the slice direction SL. The coverage of the X-ray detector 32 is 40 mm, for example.

  The collimator device 33 includes a plurality of collimator plates 33P that are erected so as to divide the plurality of X-ray detection elements 32i in at least the channel direction CH. That is, the plurality of collimator plates 33P are disposed so as to be positioned at least at the boundary between the X-ray detection elements 32i adjacent to each other in the channel direction CH. The plate surface of the collimator plate 33P is adjusted to be parallel to the direction in which the X-ray 81 is irradiated from the X-ray focal point 31F of the X-ray tube 31 and the X-ray irradiation direction E. The collimator plate 33P is made of a heavy metal having X-ray absorption, such as tungsten or molybdenum. The plate thickness Δt of the collimator plate 33P is, for example, 0.2 mm.

  Note that the X-ray detector 32 may be configured by modularizing a predetermined number of X-ray detection elements 32i and arranging a plurality of detector modules. Similarly, the collimator device 33 may be configured by modularizing a predetermined number of collimator plates 33P and arranging the collimator modules 33P.

  FIG. 3 is an enlarged cross-sectional view showing the configuration of the X-ray detection element 32i. This figure is a view when a cross section of the X-ray detection element 32i parallel to the channel direction CH is viewed in the slice direction SL.

  The X-ray detection element 32 i is provided on a ceramic substrate 51, a light conversion element 52, a first photoelectric conversion element 53, a second photoelectric conversion element 54, a light reflection plate 55, and a lid portion 56. And have.

  The light conversion element 52 is an element that converts incident X-rays 81 into photons 82 and emits them. The X-ray 81 usually has a predetermined X-ray energy distribution. X-rays having a specific X-ray energy can be considered as a mass of X-ray particles according to the magnitude of the X-ray energy. The light conversion element 52 converts the X-ray particles into photons 82 with a predetermined probability while maintaining the mass of the X-ray particles. That is, when the X-rays 81 are incident, the light conversion element 52 emits a number of photons 82 corresponding to the X-ray energy at substantially the same time for each X-ray energy.

  The light conversion element 52 is, for example, a scintillator. The light conversion element 52 has a substantially rectangular shape or cubic shape. The light conversion element 52 has an X-ray incident surface substantially perpendicular to the X-ray irradiation direction, that is, the X-ray incident direction E, and side surfaces parallel to the X-ray incident direction E are channel directions CH and They are arranged so as to be substantially parallel to the slice direction SL.

  The first photoelectric conversion element 53 is an element that converts the incident photon 82 into an electrical signal and outputs the accumulated signal G1. The accumulated signal G1 is a signal obtained by integrating the intensity of the electric signal corresponding to each photon incident within a certain time in the time axis direction. The first photoelectric conversion element 53 is, for example, a photodiode. The first photoelectric conversion element 53 has a substantially plate-like rectangular shape, and the shape of the plate surface approximates the shape of the X-ray emission surface of the light conversion element 52. The first photoelectric conversion element 53 is disposed on the X-ray emission surface side of the light conversion element 52, that is, on the bottom surface side when the X-ray incident direction is the vertical direction. Further, the first photoelectric conversion element 53 is arranged such that the plate thickness direction is substantially parallel to the X-ray incident direction E, and the edge of the plate surface is substantially parallel to the channel direction CH and the slice direction SL. ing. With such a configuration, the first photoelectric conversion element 53 receives the photons 82 emitted from the light conversion element 52 and outputs an electrical signal having an intensity corresponding to the number of photons 82 received within a predetermined time.

  The second photoelectric conversion element 54 converts the incident photon 82 into an electric signal and outputs the independent signal G2. The independent signal G2 is an electrical pulse signal corresponding to each incident photon 82. When the photon 82 group is incident simultaneously, the second photoelectric conversion element 52 outputs a pulse signal having a wave height corresponding to the number of photons 82 constituting the photon 82 group. The second photoelectric conversion element 54 is a semiconductor device (semi-conductor device) suitable for so-called photon counting, and is, for example, a silicon-multiplier (Si-MP). As one type of Si-MP, for example, there is a Multi-Pixel Photon Counter (MPPC) (registered trademark) series provided by Hamamatsu Photonics. The second photoelectric conversion element 54 has a substantially plate-like rectangular shape, and the shape of the plate surface approximates the shape of the side surface of the light conversion element 52. The second photoelectric conversion element 54 is disposed on the side of the light conversion element 52 that is orthogonal to the X-ray incident direction, that is, on the side surface when the X-ray incident direction is the vertical direction. With such a configuration, the second photoelectric conversion element 54 receives the photons 82 emitted from the light conversion element 52 and outputs an electric signal in a pulse shape. When the intensity of transmitted X-rays of the subject 91 is sufficiently low, the photon 82 group for each X-ray energy is emitted from the light conversion element 52 while being scattered in the time axis direction. At this time, the second photoelectric conversion element 54 generates a pulse signal having a pulse height corresponding to the magnitude of the X-ray energy for each X-ray energy according to the dose of X-rays having the X-ray energy. The output is output in a state of being scattered in the time axis direction by the number. Therefore, if the pulse signals output within a predetermined time are counted for each wave height, the dose of the transmitted X-rays of the subject 91 can be known for each X-ray energy. Further, if all the pulse signals output within a certain time are counted regardless of the wave height, the dose of the entire transmitted X-ray of the subject 91 can be known.

  A conductive pattern (not shown) formed on the ceramic substrate 51 is connected to each of the first and second photoelectric conversion elements 53 and 54, and an accumulated signal from the first photoelectric conversion element 53 is connected to the first and second photoelectric conversion elements 53 and 54. The independent signal G2 from G1 and the second photoelectric conversion element 54 is output to the outside through the conductor pattern.

  Note that only one second photoelectric conversion element 54 may be disposed for one side surface for each light conversion element 52, but a plurality of side surfaces are provided for each light conversion element 52. May be arranged one by one. For example, one light conversion element 52 may be disposed on one side surface in the channel direction CH and one side surface in the slice direction SL.

  Further, among the four side surfaces of the light conversion element 52, a member that does not transmit light is provided on the side surface on which the second photoelectric conversion element 54 is not disposed. Here, as such a member, a light reflection plate 55 is provided so that photons emitted from the light conversion element 52 are reflected.

  Further, on the X-ray incident surface side of the light conversion element 52, a cover part 56 having X-ray transparency may be provided in order to protect the light conversion element 52. The lid part 56 is made of plastic resin, for example.

  In the present embodiment, the accumulated signal G1 output from the first photoelectric conversion element 53 is used for collecting projection data according to the conventional method, and the independent signal G2 output from the second photoelectric conversion element 54 is photon counting. Used to collect projection data by the method.

  In the conventional method, the X-ray detection sensitivity is relatively low. Therefore, in order to obtain a signal with a satisfactory S / N ratio, it is necessary to earn photons 82 emitted from the light conversion element 52 to some extent, and the subject 91 must be irradiated with a relatively high dose of X-rays.

  On the other hand, in the photon counting method, the X-ray detection sensitivity is very high. That is, even when the dose of the X-ray 81 is low and the number of photons 82 emitted from the light conversion element 52 is small, a signal with a sufficient SN ratio can be obtained. However, if the dose of the X-ray 81 is large, a phenomenon called so-called pile up occurs, and the signal cannot be resolved in the time axis direction, so it cannot be returned. Therefore, in the photon counting method, the subject 91 must be irradiated with a relatively very low dose of X-rays 81.

  In this example, a current mode, a pulse mode, and a current & pulse mode are prepared as shooting modes. The shooting mode selection unit 41 selects any one of these modes based on the operation of the operator.

  When the current mode is selected, the imaging control unit 44 performs X-ray imaging in order to perform imaging when the tube current of the X-ray tube 31 is the first tube current value A1 corresponding to the first dose D1. Each part including the pipe 31 and the like is controlled. By such control, imaging is performed with the X-ray of the first dose D1, and projection data of a plurality of views with the first dose D1 is collected.

  When the pulse mode is selected, the imaging control unit 44 has the tube current of the X-ray tube 31 lower than the first tube current value A1 and the second tube current value A2 corresponding to the second dose D2. Each part including the X-ray tube 31 and the like is controlled so as to perform imaging. By such control, imaging is performed with the X-ray of the second dose D2, and projection data of a plurality of views with the second dose D2 is collected.

  When the current & pulse mode is selected, the imaging control unit 44 performs imaging when the tube current of the X-ray tube 31 is the first tube current value A1 corresponding to the first dose D1, X-rays In order to perform imaging when the tube current of the tube 31 is lower than the first tube current value A1 and the second tube current value A2 corresponding to the second dose D2, the respective units including the X-ray tube 31 and the like are controlled. To do. By such control, imaging is performed with the X-ray of the first dose D1 and the X-ray of the second dose D2 lower than the first dose D1, and projection data of a plurality of views is collected for each dose. The

  FIG. 4 is a diagram showing an example of a time chart of tube current at the time of photographing. This figure is an example in the current & pulse mode.

  In this example, when the current & pulse mode is selected, the tube current of the X-ray tube 31 is changed into the first tube current value A1 and the second tube current value A2 for each view as shown in FIG. The subject 91 is irradiated with X-rays 81 while switching to the above. For the view, for example, one gantry rotation time is about 1 second, and about 1000 views are assigned to one gantry rotation.

  The X-ray detector 32 receives the transmitted X-rays of the subject 91 for each view and outputs a signal. The data collection unit 34 receives a signal in the current mode for each time width when the tube current is the first tube current value A1 in the time width corresponding to one view. That is, the accumulation signal G1 from each first photoelectric conversion element 53 is received. Then, based on the intensity of the accumulated signal G1, projection data of the view is determined. The data collection unit receives a signal in the pulse mode for each time width when the tube current is the second tube current value A2 in the time width corresponding to one view. That is, the independent signal G2 from each second photoelectric conversion element 54 is received. Then, based on the count value of the independent signal G2, projection data of the view is determined.

  The data collecting unit 34 can also count the independent signal G2 for each wave height when counting the independent signal from the second photoelectric conversion element 53. When the crest value is plotted on the horizontal axis and the count value is plotted on the vertical axis, an X-ray energy spectrum representing the transmitted X-ray dose of the subject 91 for each X-ray energy can be obtained. Spectral imaging can be performed by extracting only transmitted X-ray intensity of specific X-ray energy, determining projection data, and reconstructing an image based on the projection data. Spectral imaging makes it possible to emphasize and draw a substance that absorbs specific X-ray energy well, so that pathological functional imaging such as blood flow (perfusion) of the subject 91 can be realized.

  By the way, regarding the data collection method, the conventional method using the accumulated signal G1 and the photon counting method (PC method) using the independent signal G2 each have the characteristics shown in the following table.

  That is, in the conventional method using the accumulated signal G1, the dynamic range of the X-ray detection signal can be widened, and the reconstructed image can be made familiar, but pathological functional imaging is difficult. The low exposure of the subject 91 and the signal noise ratio of the X-ray detection signal are bad.

  On the other hand, in the photon counting method using the independent signal G2, pathological functional imaging is possible, and the low exposure of the subject and the SN ratio of the X-ray detection signal are good. Since the pile-up occurs when the dose of X-rays is large, the dynamic range of the X-ray detection signal is narrow, and the depiction may differ from the conventional depiction by the amount of good sensitivity, and some operators 92 may feel uncomfortable.

  Thus, for example, when it is desired to perform shape imaging of the subject 91 with a sense of incongruity, the current mode according to the conventional method is used. For example, when it is desired to perform pathological functional imaging of the subject 91 and when it is desired to perform imaging with low exposure without having to worry too much about the dynamic range, a pulse mode based on the photon counting method is used. For example, when it is desired to display an image representing the shape of the subject 91 and an image representing a pathological function (perfusion, etc.) in an overlapping manner, the current & pulse mode is used.

  Hereafter, the flow of operation in the X-ray CT apparatus according to the present embodiment will be described.

  FIG. 5 is a flowchart showing an example of the operation flow in the X-ray CT apparatus according to the present embodiment.

  In step S1, a shooting mode is selected. Specifically, the shooting mode selection unit 41 of the operation console 4 selects any of the current mode, the pulse mode, and the current & pulse mode as the shooting mode based on the operation of the operator 92.

  In step S2, shooting conditions are set. Specifically, the imaging condition setting unit 42 of the operation console 4 determines the tube voltage and tube current of the X-ray tube 31, the imaging range of the subject 91, the gantry rotation speed, the slice thickness, and the reconstruction based on the operation of the operator. Set functions, etc.

  In step S3, data collection is set. Specifically, the data collection unit 34 performs setting so as to generate and collect either one or both of projection data based on the accumulated signal and projection data based on the independent signal according to the selected photographing mode.

  In step S4, the subject 91 is imaged. Specifically, the imaging control unit 44 of the operation console 4 controls each unit according to the set imaging conditions and images the subject 91. The data collection unit 34 collects projection data. At this time, when the pulse mode or the current & pulse mode is selected as the imaging mode, projection data based on the independent signal is generated for each X-ray energy.

  In step S5, the image is reconstructed. Specifically, the image reconstruction unit 45 of the operation console 4 reconstructs an image by applying a reconstruction function to the collected projection data. At this time, when the current mode or the current & pulse mode is selected, the image is reconstructed using the projection data based on the accumulated signal G1. In addition, when the pulse mode or the current & pulse mode is selected, the projection data for each X-ray energy based on the independent signal G2 is used to regenerate an image for each X-ray energy and an image corresponding to the entire X-ray energy. Configure.

  In step S6, an image is displayed. Specifically, the display control unit 46 of the operation console 4 displays the reconstructed image on a monitor screen. At this time, when the current mode or the current & pulse mode is selected, an image based on the accumulation signal G1 is displayed. When the pulse mode or the current & pulse mode is selected, an image based on the independent signal G2 is displayed. At this time, the operator 92 can input desired X-ray energy. The display control unit 46 can display an image corresponding to desired X-ray energy based on this input information. Of course, an image corresponding to the entire X-ray energy can also be displayed. The image based on the accumulated signal G1 and the image based on the independent signal G2 can be displayed in an overlapping manner after the subject included in the image is aligned.

  According to the present embodiment, the X-ray detection element 32 i constituting the X-ray detector 32 is converted by the light conversion element 52 that converts the X-ray 81 into the photon 82 and the light conversion element 52 and is incident thereon. A first photoelectric conversion element 52 having a current mode structure that converts a photon 82 into an electrical accumulated signal G1 and outputs it, and a photon 82 converted and incident by the light conversion element 52 into an electrical independent signal G2 The second photoelectric conversion element 54 having a pulse mode structure that outputs in the same manner, so that the merits of the pulse mode are maintained as necessary while maintaining the image quality and work flow in the current mode with proven results. (Merit) can be enjoyed.

  In the present embodiment, the second photoelectric conversion element 54 is disposed on the side surface of the light conversion element 52, and also serves as a conventional partition plate provided between adjacent light conversion elements 52. The shape and size of the light conversion element 52 are hardly adversely affected. That is, it is possible to reduce design changes such as the mechanical structure of the photographing apparatus and the algorithm for image reconstruction (algorithm), and to suppress an increase in production cost and production man-hour.

  The embodiment of the invention is not limited to the above, and various modifications can be made without departing from the spirit of the invention.

  For example, in the above embodiment, the first photoelectric conversion element 53 is provided on the bottom surface of the light conversion element 52, and the second photoelectric conversion element 54 is provided on the side surface of the light conversion element 52. Without being limited thereto, the first photoelectric conversion element 53 is provided on the first light emission surface of the light conversion element 52, and the second photoelectric conversion element 54 is the first photoelectric conversion element 52 in the light conversion element 52. The second light emission surface may be different from the light emission surface. However, the first photoelectric conversion element 53 that outputs an accumulation signal is provided on the bottom surface of the light conversion element 52, and the second photoelectric conversion element 54 that outputs an independent signal is provided on the side surface of the light conversion element 52. The detector structure is close to the current one. For this reason, this structure is more preferable because it facilitates design, contributes to cost and manufacturing man-hour reduction, and stability of performance.

  Further, for example, a detector having the same structure as the detector in the above-described embodiment can be applied to a radiographic apparatus such as a general imaging apparatus that performs radiographic imaging of the chest and the like, and such a radiographic apparatus is also an invention. It is one Embodiment.

DESCRIPTION OF SYMBOLS 1 X-ray CT apparatus 2 Imaging table 3 Scanning gantry 3B Cavity part 31 X-ray tube 32 X-ray detector 32i X-ray detection element 33 Collimator apparatus 34 Data collection part 4 Operation console 41 Imaging mode selection part 43 Imaging condition setting part 44 Imaging | photography Control unit 45 Image reconstruction unit 46 Display control unit 51 Ceramic substrate 52 Light conversion element 53 First photoelectric conversion element 54 Second photoelectric conversion element 55 Light reflection plate 56 Lid 81 X-ray 82 Photon 91 Subject 92 Operator ISO Isocenter

Claims (9)

A radiation detector in which a plurality of radiation detection elements are arranged in at least one direction,
Each of the plurality of radiation detection elements includes:
A single-layer light conversion element that converts incident radiation into photons;
A first photoelectric conversion element that is provided on one of a bottom surface and a side surface of the light conversion element, converts the incident photons into an electrical accumulated signal, and outputs the electrical accumulated signal;
A second photoelectric conversion element that is provided on the other of the bottom surface and the side surface of the light conversion element, converts the incident photons into an electrical independent signal, and outputs the electrical independent signal;
Including radiation detector. The light conversion element includes a scintillator,
The first photoelectric conversion element includes a photodiode,
The radiation detector according to claim 1, wherein the second photoelectric conversion element includes Si-PM (Silicon Photomultiplier). The radiation imaging apparatus comprising a radiation detector according to claim 1 or 2. First determination means for determining the intensity of a radiation detection signal in the radiation detection element based on the magnitude of the accumulated signal output from the first photoelectric conversion element included in the radiation detection element;
A second determination that determines the intensity of a radiation detection signal in the radiation detection element based on a count value within a predetermined time of the independent signal output from the second photoelectric conversion element included in the radiation detection element. And a radiation imaging apparatus according to claim 3 . First reconstruction means for reconstructing a first image based on first projection data based on the intensity of the radiation detection signal determined by the first determination means;
Of claim 4 and a second reconstruction means for reconstructing a second image based on the second projection data by the intensity of the detection signal of the radiation determined by the second determining means Radiography equipment. The second determining means determines the intensity of a detection signal of radiation having specific energy in the radiation detection element based on a count value within a certain time of a signal having a specific wave height among the independent signals,
The radiation imaging apparatus according to claim 5 , wherein the second reconstruction unit reconstructs a second image based on projection data based on an intensity of a detection signal of radiation having the specific energy. The radiographic apparatus according to claim 5 or claim 6 further comprising a display means for displaying overlapping and said second image and the first image. Control means for controlling the radiation source so as to irradiate and image radiation with a first dose and radiation with a second dose lower than the first dose;
The first determining means determines the intensity of the radiation detection signal based on the accumulated signal at the time of radiation irradiation with the first dose,
Said second determining means determines the intensity of the detection signal of the radiation on the basis of the independent signals during irradiation by the second dose, according to any one of claims 7 claims 4 Radiography equipment. The imaging is radiation tomography,
The radiographic apparatus according to claim 8 , wherein the control unit controls the radiation dose to be irradiated to switch between the first dose and the second dose for each view.

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