JP2011055909A - X-ray detector for x-ray computer tomography device and x-ray computer tomography device - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To discriminate X-rays which have transmitted through a subject for each energy by a simple structure. <P>SOLUTION: The X-ray detector includes a plurality of X-ray detection elements 10 arrayed along a first direction, each of the X-ray detection elements 10 includes: a plurality of scintillator parts 41, 42 which are laminated along a second direction vertical to the first direction, and of which the wavelengths indicating a sensitivity peak to the energy of incident X-rays are different; and a plurality of light receiving parts 31, 32 which are arrayed along the first direction on the back surface of the scintillator part 42, and of which the wavelengths indicating the sensitivity peak to incident light are different. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to an X-ray detector capable of discriminating incident X-ray energy and an X-ray computed tomography apparatus having an X-ray detector capable of discriminating incident X-ray energy.

  In an X-ray computed tomography apparatus (hereinafter referred to as CT), X-rays emitted from an X-ray generator pass through a subject. The X-ray detector detects the X-ray intensity attenuated according to the X-ray attenuation coefficient of the substance in the transmitted object. The X-ray attenuation coefficient of the substance in the subject changes according to the energy of the transmitted X-ray. When the X-ray energy is low, the X-ray attenuation coefficient is large. Conversely, when the X-ray energy is high, the X-ray attenuation coefficient is small. This is because, when the reconstructed image shows a close CT value and the difference in the substance cannot be discriminated, the reconstructed image is obtained by discriminating the energy of the detected X-ray, thereby showing the different CT value. It means that it can be discriminated.

  In order to obtain a CT image by discriminating the energy of X-rays transmitted through the subject, a conventional method has been devised or commercialized by changing the applied voltage to the X-ray tube of the X-ray generation unit. .

  There are two methods for scanning by changing the applied voltage to the X-ray tube: a method of changing the applied voltage for each sampling point, and a method of changing the applied voltage for each rotation. The former requires a technique for changing a high voltage and a large current at a high speed and a technique for sampling all data at the same timing as this, and there is a problem that the equipment cost increases. The latter has a problem that it is difficult to compare images taken with different energies because the sampling periods of data with different X-ray energies are not the same.

  An object of the present invention is to discriminate, for each energy, X-rays transmitted through a subject with a simple structure in an X-ray detector for an X-ray computed tomography apparatus and an X-ray computed tomography apparatus.

  In order to achieve the above object, the present invention takes the following measures.

  An X-ray detector according to a first aspect of the present invention has a plurality of X-ray detection elements arranged along a first direction, and each X-ray detection element is a second perpendicular to the first direction. A plurality of scintillator portions that are stacked along the direction and have different wavelengths indicating sensitivity peaks with respect to the energy of incident X-rays, and wavelengths that are arranged along the first direction on the back surface of the scintillator portions and indicate sensitivity peaks with respect to incident light And a plurality of light receiving portions different from each other.

  An X-ray computed tomography apparatus according to a second aspect of the present invention detects an X-ray generated from an X-ray generator and an X-ray generator that generates X-rays and transmitted through the subject in a first direction. An X-ray detector comprising: an X-ray detector having a plurality of X-ray detection elements arranged along; and a reconstruction processing unit for reconstructing at least one image based on an output from the X-ray detector Each element is stacked along a second direction perpendicular to the first direction, and includes a plurality of scintillator portions having different wavelengths indicating sensitivity peaks with respect to energy of incident X-rays, and a first direction on the back surface of the scintillator portion. And a plurality of light receiving portions having different wavelengths indicating sensitivity peaks with respect to incident light.

  According to the present invention, X-rays transmitted through a subject with a simple structure can be discriminated for each energy.

1 is a cross-sectional view illustrating an example of a configuration of a one-channel X-ray detection element 10 in an X-ray detector according to a first embodiment of the present invention. FIG. 2 is a diagram illustrating the definition of the first direction and the second direction described in FIG. 1. FIG. 3 shows an example of the light emission characteristics of the first scintillator portion 41 and the second scintillator portion 42 shown in FIG. 1, with the maximum sensitivity of the scintillator portions 41 and 42 with respect to the energy of X-rays incident on the scintillator 11. It is a figure which shows the peak wavelength of the fluorescence emitted from the scintillator parts 41 and 42. FIG. FIG. 4 shows, as an example of signal conversion characteristics in the first photodiode 21 and the second photodiode 22 which are the light detection elements shown in FIG. 1, an electric signal with respect to the wavelength of fluorescence incident on each photodiode. It is a figure which shows the sensitivity peak of output. FIG. 5 shows the output from the first photodiode 21 and the second photodiode 22 regarding the cross section of the configuration of the X-ray detection element 10 for one channel in the X-ray detector according to the first embodiment. It is a figure which shows an example integrated as one output. FIG. 6 is a diagram showing a configuration of an X-ray computed tomography apparatus according to the second embodiment.

Hereinafter, embodiments of the present invention will be described with reference to the drawings.
(First embodiment)
FIG. 1 is a diagram showing a cross section of an example of the configuration of the X-ray detector 10 for one channel in the X-ray detector according to the present embodiment.
The first direction is a direction in which the X-ray detection elements 10 are arranged, and the second direction is a direction perpendicular to the first direction. As an example, the first direction and the second direction will be described with reference to FIG. 2 which is a cross-sectional view of the gantry 200 of the X-ray computed tomography apparatus.

  The gantry 200 houses a rotation support mechanism. The rotation support mechanism includes a rotation ring 102, a ring support mechanism that supports the rotation ring 102 so as to be rotatable about the rotation axis Z, and a drive unit 107 (electric motor) that drives the rotation of the ring. The X axis is a straight line that is orthogonal to the rotation axis Z and passes through the focal point 100 of the emitted X-ray. The Y axis is a straight line orthogonal to the X axis and the rotation axis Z. For convenience of explanation, the XYZ coordinate system will be described as a rotating coordinate system that rotates about the rotation axis Z.

  In the X-ray generator 101, X-rays emitted from the focal point 100 of the X-rays emitted are, for example, cone-beam (pyramidal) by a collimator unit 118 attached to the X-ray emission window of the X-ray generator 101. ). The X-ray radiation range is indicated by a dotted line.

  An X-ray generator 101 and an X-ray detector 203 are mounted on the rotating ring 102. The X-ray detector 203 is attached at a position and an angle facing the X-ray generation unit 101 with the rotation axis Z interposed therebetween.

  The X-ray detector 203 has a plurality of X-ray detection elements. Here, it is assumed that a single X-ray detection element constitutes a single channel. The plurality of channels are orthogonal to the rotation axis Z and have an arc direction whose radius is a distance from the center to the center of the light receiving portion of the X-ray detection element 10 for one channel with the X-ray focal point 100 as the center. They are arranged in one dimension with respect to (channel direction). This channel direction is the first direction. The direction perpendicular to the first direction and the direction parallel to the rotation axis Z (slice direction) is the second direction. That is, the second direction is a direction passing through the center of the light receiving portion of each X-ray detection element and the X-ray focal point 100, and is substantially the radial direction of a circle centering on the X-ray focal point 100. Further, the X-ray detector 203 may be configured by a plurality of modules in which a plurality of X-ray detection elements are arranged in a line. Each module is arranged in a one-dimensional manner in a substantially arc direction along the channel direction. At this time, the first direction is a direction in which a plurality of modules are arranged in a row, and is substantially equivalent to the first direction. The second direction is a direction passing through the X-ray focal point 100 and the center of each module, and is orthogonal to the first direction and the rotation axis Z.

  The plurality of X-ray detection elements may be two-dimensionally arranged in two directions, that is, a channel direction and a slice direction. That is, the two-dimensional arrangement is configured by arranging a plurality of channels arranged in a one-dimensional manner along the channel direction in a plurality of rows in the slice direction. An area detector having such a two-dimensional X-ray detection element array may be configured by arranging a plurality of the above-described modules arranged in a one-dimensional shape in a substantially arc direction in a plurality of rows in the slice direction.

  The area detector may be composed of a plurality of modules. Each module has a plurality of X-ray detection elements arranged in an M × N matrix. The plurality of modules are arranged two-dimensionally in the channel direction and the slice direction.

Returning to FIG. 1, the configuration of the X-ray detection element 10 for one channel will be described. The front surface is a surface closer to the X-ray generation unit 101. The back surface is a surface far from the X-ray generation unit 101. The X-ray detection element 10 includes a scintillator 11 and a light receiving unit 12. A scintillator 11 is disposed on the front surface of the light receiving unit 12. The scintillator 11 has a sensitivity peak with respect to the first energy e ₁ and the second energy e ₂ (e ₁ <e ₂ ). The scintillator 11 includes a first scintillator portion 41 and a second scintillator portion 42. First scintillator portion 41 has a first sensitivity peak relative energy e _1. Second scintillator portion 42 has a second sensitivity peak relative energy e _2. The first scintillator portion 41 is overlaid on the front surface of the second scintillator portion 42.

The light receiving unit 12 includes a first light receiving portion 31 and a second light receiving portion 32. The first light receiving portion 31 has a sensitivity peak with respect to the first wavelength. The first wavelength is equivalent to the wavelength of the fluorescence corresponding to the X-ray component of the first energy e ₁ in the first scintillator portion 41. The second light receiving portion 32 has a sensitivity peak with respect to the second wavelength. The second wavelength is equivalent to the wavelength of the fluorescence corresponding to the X-ray component of the second energy e ₂ in the second scintillator portion 42. The first light receiving portion 31 and the second light receiving portion 32 are juxtaposed on the back surface of the second scintillator portion 42.

  Then, the structure of the 1st light-receiving part 31 and the 2nd light-receiving part 32 is demonstrated. In FIG. 1, a photodiode is used as an example to detect fluorescence emitted from the scintillator portions 41 and 42. The first light receiving portion 31 includes a first photodiode 21 and a first optical filter 24. The first optical filter 24 is disposed on the front surface of the first photodiode 21. The first photodiode 21 has a sensitivity peak with respect to the first wavelength. The first optical filter 24 has a characteristic that the transmittance (attenuation rate) of light in a specific wavelength region is higher (lower) than the transmittance (attenuation rate) in other wavelength regions. The specific wavelength region is a wavelength region having a predetermined width with the first wavelength being substantially the center. For example, the width is designed to be an absolute value obtained by subtracting the second wavelength from the first wavelength.

  The first optical filter 24 may be a filter having a characteristic that the transmittance (attenuation rate) of light of a specific wavelength or higher is higher (lower) than the transmittance (attenuation rate) of less than the specific wavelength. The specific wavelength is, for example, an average wavelength of the first wavelength and the second wavelength. In this case, in order to obtain energy discrimination, the second optical filter 25 has a transmittance (attenuation rate) of light less than a specific wavelength higher (lower) than a transmittance (attenuation rate) of the specific wavelength or more. A filter having characteristics is employed.

  The second light receiving portion 32 includes a second photodiode 22 and a second optical filter 25. The second optical filter 25 is disposed on the front surface of the second photodiode 22. The second photodiode 22 has a sensitivity peak with respect to the second wavelength. The second optical filter 25 has a characteristic that the transmittance (attenuation rate) of light in a specific wavelength region is higher (lower) than the transmittance (attenuation rate) in other wavelength regions. The specific wavelength region is a wavelength region having a predetermined width that is substantially centered on the second wavelength. For example, the width is designed to be an absolute value obtained by subtracting the second wavelength from the first wavelength.

  In the configuration of each of the light receiving portions 31 and 32, the wavelength of the fluorescence corresponding to the X-ray component of the energy emitted from the scintillator portions 41 and 42, that is, the first wavelength and the second wavelength is obtained by combining an optical filter with a photodiode. By discriminating with an optical filter, it is possible to improve the discriminability of discriminating X-rays by energy rather than discriminating by photodiodes 21 and 22 alone.

FIG. 3 shows an example of the light emission characteristics of the scintillator portions 41 and 42. The peak wavelength means a wavelength λ of light emitted by an energy component of X-rays showing a so-called sensitivity peak showing the highest value of conversion efficiency from X-rays to light. Corresponding to the sensitivity peaks of the scintillator portions 41 and 42 with respect to X-ray energy, the peak wavelengths of the fluorescence emitted from the scintillator portions 41 and 42 are shown. For example, the first scintillator portion 41 has a sensitivity peak with respect to X-rays of the first energy e ₁ . Peak wavelength at this time (the first wavelength) is lambda _1. Second scintillator portion 42 has a peak sensitivity to the second X-ray energy e _2. Peak wavelength at this time (the second wavelength) is lambda _2. The peak wavelength λ ₁ is longer than the peak wavelength λ ₂ and the second energy e ₂ is higher than the first energy e ₁ . As an example, a yellow transparent scintillator that emits light having a wavelength of 512 nm is known as the peak wavelength λ ₁ . The peak wavelength lambda _2, the scintillator colorless transparent are known to emit light having a wavelength of 480 nm. In addition, due to the colors of these scintillators, the second scintillator 42 is more transparent to visible light than the first scintillator 41.

  In this manner, by combining the scintillator portions 41 and 42 having different fluorescence peak wavelengths with the light receiving unit 12 in accordance with the sensitivity peak with respect to the energy of X-rays, the X-rays are discriminated with high discrimination for each energy. Can do.

Next, the first scintillator portion 41 in the scintillator 11 is arranged on the front surface of the second scintillator portion 42, thereby providing the following effects. In the scintillator 11, the second having a sensitivity peak for the high-energy e ₂ scintillator portion 42 in front of the light receiving portion 12, the first having a peak sensitivity for low-energy e ₁ of the scintillator portion 41 second By disposing in front of the scintillator portion 42, the loss of X-ray energy can be suppressed as compared with the case of disposing in the reverse order. For this purpose, the scintillator portions 41 and 42 are superposed on the X-ray focal point 100 immediately before the light receiving unit 12 in order of increasing energy indicating a sensitivity peak.

  In addition, the following effects are also obtained by arranging the first scintillator portion 41 on the front surface of the second scintillator portion. When the fluorescence emitted from the first scintillator portion 41 passes through the second scintillator portion 42 before reaching the light receiving unit 12, the transparency of the visible light in the second scintillator portion 42 that is transmitted is low. The second scintillator portion 42 (colorless and transparent) is more transparent to visible light than the first scintillator portion 41 (transparent yellow), so the first scintillator portion 41 is absorbed. Is arranged on the front surface of the second scintillator portion 42, so that the fluorescence detection efficiency can be improved as compared with the case where the second scintillator portion 42 is arranged in the reverse order.

  In the present embodiment in FIG. 1, the scintillator 11 is disposed on the front surface of the light receiving unit 12, but scintillators having different light emission characteristics such as the first scintillator portion 41 and the second scintillator portion 42 shown in FIG. A mixture of these may be arranged on the front surface of the light receiving unit 12.

FIG. 4 shows, as an example of the signal conversion characteristics of the photodiode, the output of the first photodiode 21 and the second photodiode 22 by the electrical signal with respect to the wavelength of the fluorescence incident on the photodiode (light receiving wavelength). Yes. The first photodiode 21 has an output peak value with respect to the light receiving wavelength λ ₁ . The light receiving wavelength λ ₁ is also a peak wavelength in the first scintillator portion 41. The second photodiode 22 has a peak value of the output to the light-receiving wavelength lambda _2. The light receiving wavelength λ ₂ is also the peak wavelength of the second scintillator portion 42.

Thus, the fluorescence incident on the first photodiode 21 and the second photodiode 22 can be distinguished for each received wavelength. Further, by combining the outputs from the photodiodes 21 and 22 with separate wirings and combining the photodiodes 21 and 22 having different light receiving wavelengths with respect to the output peaks with the scintillator portions 41 and 42, the X-rays are output on separate wirings for each energy. In the above description, a photodiode is used as the light detection element in the light receiving unit 12, but other light detection elements may be used.

(function)
Hereinafter, with reference to FIG. 1, in the X-ray detector of the present embodiment, X-rays incident on the X-ray detection element 10 for one channel are discriminated based on the peak wavelength of fluorescence emitted from the scintillator 11 for each energy. The function to perform will be described.

X-rays having a certain energy spectrum are incident on the first scintillator portion 41 of the one-channel X-ray detection element 10. The first scintillator portion 41 emits fluorescence having a peak wavelength λ ₁ corresponding to incident X-ray energy e ₁ in accordance with the emission characteristics shown in FIG. In FIG. 1, a place where fluorescence is emitted by incident X-rays is indicated by 1. The fluorescence having the peak wavelength λ ₁ has a characteristic in which the transmittance in the wavelength region having a width of | λ ₁ −λ ₂ | with the peak wavelength λ ₁ substantially in the center is higher than the transmittance in other wavelength regions. The light passes through the optical filter 24 and enters the first photodiode 21. The first photodiode 21 converts the incident fluorescent light having the peak wavelength λ ₁ into an electrical signal according to the signal conversion characteristics shown in FIG. The fluorescence having the peak wavelength λ ₁ has a characteristic that the transmittance in the wavelength region having a width of | λ ₂ −λ ₁ | with the peak wavelength λ ₂ as the center is higher than the transmittance in the other wavelength regions. Since it is blocked by the second optical filter 25, it does not enter the second photodiode 22.

X-rays having an energy spectrum that has not been converted into fluorescence by the first scintillator portion 41 enter the second scintillator portion 42. The second scintillator portion 42 emits fluorescence having a peak wavelength λ ₂ corresponding to the incident energy e ₂ of the X-ray in accordance with the light emission characteristics shown in FIG. In FIG. 1, a place where fluorescence is emitted by incident X-rays is indicated by 2. The fluorescence having the peak wavelength λ ₂ passes through the second optical filter 25 and enters the second photodiode 22. The second photodiode 22 converts the incident fluorescence having the peak wavelength λ ₂ into an electrical signal according to the signal conversion characteristics shown in FIG. Note that the fluorescence having the peak wavelength λ ₂ is blocked by the first optical filter 24 and therefore does not enter the first photodiode 21.

  The function of discriminating X-rays for each energy in the X-rays incident on the one-channel X-ray detection element 10 has been described above. For incident X-rays, two types of output are obtained: an electrical signal output by low energy X-rays and an electrical signal output by high energy X-rays. The wiring shown in FIG. 1 is configured to extract these two types of outputs from the light receiving portions 31 and 32 by separate wiring in the X-ray detection element 10 for one channel. With this wiring, the X-rays incident on the X-ray detector can be distinguished for each energy. Note that the wiring obtained by integrating these two types of outputs is the wiring shown in FIG. With this wiring, the output obtained by discriminating X-rays for each energy can be integrated.

  According to the configuration described above, the following effects can be obtained.

  According to the X-ray detection element 10 of each channel in the present X-ray detector, the difference in the fluorescence peak wavelength is obtained by using the scintillator portions 41 and 42 having different fluorescence peak wavelengths according to the sensitivity peak to the X-ray energy. X-rays can be discriminated for each energy. Subsequently, the incident fluorescence peak wavelength can be distinguished by the photodiodes 21 and 22 that output electrical signals based on the fluorescence peak wavelength. From these facts, X-rays incident on the X-ray detection elements of the respective channels can be distinguished for each energy. Further, by combining the photodiodes 21 and 22 with the optical filters 24 and 25, the optical wavelength of the fluorescence corresponding to the X-ray component of the energy emitted from the scintillator portions 41 and 42, that is, the first wavelength and the second wavelength can be changed. By discriminating with 24 and 25, it is possible to improve the discriminability of discriminating X-rays with energy, rather than with the photodiodes 21 and 22 alone.

(Second Embodiment)
Next, a second embodiment of the present invention will be described.

  In the second embodiment described below, the same components as those in the first embodiment are denoted by the same reference numerals in order to avoid duplication of description, and detailed description thereof is omitted, and different configurations are provided. Only the elements are described.

  FIG. 6 is a diagram showing the configuration of the X-ray computed tomography apparatus according to this embodiment. The X-ray computed tomography apparatus includes a ROTATE / ROTATE-TYPE in which the X-ray generator 101 and the X-ray detector 203 are integrally rotated around the subject, and a large number of detection elements arrayed in a ring shape. There are various types, such as STATIONARY / ROTATE-TYPE, in which only the X-ray generator 101 rotates around the subject, and any type is applicable to the present embodiment. Here, it is described as ROTATE / ROTATE-TYPE. Further, in order to reconstruct an image, projection data for 360 ° around the object and projection data for 360 ° are required, and projection data for 180 ° + fan angle is also required in the half scan method. The present embodiment can be applied to any reconfiguration method. Here, the 360 ° method will be described as an example. In recent years, the so-called multi-tube type X-ray computed tomography apparatus in which a plurality of pairs of the X-ray generator 101 and the X-ray detector 203 are mounted on a rotating ring has been commercialized, and the development of peripheral technologies has been developed. Is progressing. In the present embodiment, both a conventional single-tube type and a multi-tube type are applicable. Here, a single tube type will be described.

  The X-ray generator 101 emits X-rays from the focal point 100 in response to voltage application and current supply from the high-voltage generator 109 via the slip ring 108. An X-ray generator 101 and an X-ray detector 203 are mounted on the rotating ring 102.

  As the one-channel X-ray detection element 10 in the X-ray detector 203, the one shown in FIG. 1 in the first embodiment is used. As the one-channel X-ray detection element 10 in the X-ray detector 203, the one shown in FIG. 5 in the first embodiment may be used. The X-ray detector 203 in the present embodiment is a two-dimensional area detector. An X-ray detector in which X-ray detection elements are arranged one-dimensionally may be used.

  During imaging or scanning, a subject is placed on the top 120 and inserted into a cylindrical imaging region 111 between the X-ray generator 101 and the X-ray detector 203.

  A data acquisition circuit 204 called a DAS (Data Acquisition System) is connected to the output of the X-ray detector 203. As shown in FIG. 1, the data acquisition circuit 204 outputs from the first light receiving portion 31 and from the second light receiving portion 32 based on the energy of the X-rays incident on the X-ray detection element 10 of one channel. Are connected to each channel by separate wiring according to the output. In addition, as shown in FIG. 5, you may connect the output from each light-receiving part 31 and 32 to the data collection circuit 204 for every channel as one wiring.

  The X-rays incident on the X-ray detector 203 are discriminated for each energy by connecting the light receiving portions 31 and 32 and the data acquisition circuit 204 in the X-ray detection element 10 for one channel for each channel by separate wiring. Can do. Subsequent processing is performed for each discriminated X-ray energy, except for processing in the projection data processing unit 210 and processing in the image processing unit 214.

  In addition, by connecting the outputs from the light receiving portions 31 and 32 in the X-ray detection element 10 for one channel as one wiring to the data collection circuit 204 for each channel, the X-rays incident on the X-ray detector 203 are converted into energy. The electrical signals that are discriminated for each output may be integrated for each channel and output to the data acquisition circuit 204. Since the integration of the electric signals is based on the output for each energy of the X-ray, the sensitivity of the incident X-ray can be improved as compared with the conventional X-ray detection element.

  The data acquisition circuit 204 includes an IV converter that converts a current signal of each channel of the X-ray detector 203 into a voltage, and an integration that periodically integrates the voltage signal in synchronization with an X-ray exposure cycle. A converter, an amplifier for amplifying the output signal of the integrator, and an analog / digital converter for converting the output signal of the amplifier into a digital signal are attached to each channel. Data (pure raw data) output from the data collection circuit 204 is transmitted to the preprocessing unit 106 via the non-contact data transmission unit 105 using magnetic transmission / reception or optical transmission / reception.

  The preprocessing unit 106 performs preprocessing on the pure raw data output from the data collection circuit 204. The preprocessing includes, for example, sensitivity non-uniformity correction processing between channels, X-ray strong absorber, processing for correcting signal signal drop or signal loss due to extreme metal intensity mainly. Data immediately before reconstruction processing output from the pre-processing unit 106 (raw data or projection data, here referred to as projection data) is data and a scintillator representing a view angle when data is collected. The projection data storage unit 112 including a magnetic disk, a magneto-optical disk, or a semiconductor memory is associated with X-ray energy indicating sensitivity peaks in the portions 41 and 42.

  The projection data is a set of data values corresponding to the intensity of X-rays that have passed through the subject. Here, for convenience of explanation, a set of projection data over all channels having the same view angle collected at the same time in one shot for the energy of X-rays indicating sensitivity peaks in the scintillator portions 41 and 42 is referred to as a projection data set. . In addition, the view angle is an angle in a range of 360 ° with each position of the circular orbit around which the X-ray generation unit 101 circulates around the rotation axis Z as the top of the circular orbit vertically upward from the rotation axis Z being 0 °. It is a representation. The projection data for each channel of the projection data set is identified by the view angle, cone angle, channel number, and X-ray energy indicating the maximum sensitivity in the scintillator portions 41 and 42.

  The reconstruction processing unit 114 performs the felt gunp method or the cone beam reconstruction method on the basis of the projection data set in the range of the viewing angle of 360 ° or 180 ° + fan angle for each of the discriminated X-ray energies. It has a function of reconstructing a substantially cylindrical three-dimensional image. Further, the reconstruction processing unit 114 has a function of reconstructing a two-dimensional image (tomographic image) by, for example, a fan beam reconstruction method (also referred to as a fan beam convolution back projection method) or a filtered back projection method. The Feldgump method is a reconstruction method when the projection ray intersects the reconstruction surface like a cone beam. When convolution is performed on the assumption that the cone angle is small, it is treated as a fan projection beam. Back projection is an approximate image reconstruction method that processes along a ray during scanning. The cone beam reconstruction method is a reconstruction method in which projection data is corrected in accordance with the angle of the ray with respect to the reconstruction surface, as a method of suppressing the cone angle error more than the Feldgump method.

  The reconstructed three-dimensional image or two-dimensional image is an image corresponding to each discriminated X-ray energy. Note that by applying the X-ray detection element 10 shown in FIG. 5 to the X-ray detector 203, one three-dimensional image and one two-dimensional image are reconstructed based on the output from the X-ray detector 203. The

  By reconstructing an image based on the projection data set discriminated for each X-ray energy incident on the X-ray detector 203, reconstructed images having different clinical significances can be obtained. For example, a projection data set based on the output of low-energy components of X-rays is suitable for reconstructing fat-rich images. A projection data set based on the output of the high energy component of X-rays is suitable for reconstructing an image relating to bone density.

  The projection data processing unit 210 uses a projection data set discriminated for each X-ray energy according to an input from an operator via the input unit 115 or a predetermined rule, and performs operations such as a difference between the projection data sets and composition. Process. For example, the projection data for each channel identified by the view angle, the cone angle, and the channel number and having different incident X-ray energies is subjected to arithmetic processing such as difference and synthesis, and is output to the reconstruction processing unit 114. The processing in the projection data processing unit 210 contributes to increasing the CT value contrast and improving the image quality in the reconstructed image obtained after this processing. By performing the calculation for each channel using the projection data set, the number of calculation processes is reduced as compared with the process in the image processing unit 214 using the reconstructed image described below. The influence can be reduced. Here, the predetermined rule is, for example, the rule of the arithmetic processing stored in advance in a storage unit (not shown) of the present embodiment. Without input from the operator, the host controller 110 controls the projection data processing unit 210 and performs the above arithmetic processing.

  The image processing unit 214 performs image processing such as difference and composition using the image reconstructed by the reconstruction processing unit 114 for each X-ray energy in response to an input from the operator via the input unit 115. This processing is performed by, for example, a reconstructed image corresponding to the energy of X-rays displayed on the display unit 116 (for example, a reconstructed image output from low-energy X-rays and a reconstructed image output from high-energy X-rays). Based on the input from the operator via the input unit 115. This is useful when improving the contrast of the reconstructed image or when improving the S / N ratio.

  The display unit 116 displays the image processed by the reconstruction processing unit 114 or the image processing unit 214. At this time, the display unit 116 can also display an image corresponding to each of the discriminated X-ray energies on one screen.

  According to the configuration described above, the following effects can be obtained.

  According to the X-ray computed tomography apparatus, X-rays transmitted through the subject can be distinguished for each energy. In addition, an image corresponding to the discriminated X-ray energy can be reconstructed. In addition, processing such as difference and composition can be performed using a projection data set and a reconstructed image before reconstruction processing. From these facts, the contrast resolution of the reconstructed image is improved and the difference in the substance can be discriminated in the case where the close CT value is shown on the reconstructed image and the difference in the substance cannot be discriminated.

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

  DESCRIPTION OF SYMBOLS 1 ... The place which emitted fluorescence by the incident low energy X-ray, 2 ... The place which emitted fluorescence by the incident high energy X-ray, 5 ... The fluorescence emitted from the 1st scintillator part 41, 6 ... 2nd scintillator Fluorescence emitted from the portion 42, 10 ... X-ray detection elements for one channel, 11 ... scintillator, 12 ... light receiving part, 21 ... first photodiode, 22 ... second photodiode, 24 ... first light Filter, 25 ... second optical filter, 31 ... first light receiving portion, 32 ... second light receiving portion, 41 ... first scintillator portion, 42 ... second scintillator portion, 100 ... radiation of X-rays Focus: 101 ... X-ray generator, 102 ... Rotating ring, 105 ... Non-contact data transmission unit, 106 ... Pre-processing unit, 107 ... Drive unit, 108 ... Slip ring, 109 ... High voltage generator 110: Host controller, 111: Imaging region, 112: Projection data storage unit, 114 ... Reconstruction processing unit, 115 ... Input unit, 116 ... Display unit, 118 ... Collimator unit, 120 ... Top plate, 200 ... Gantry, 203 ... X-ray detector, 204 ... Data acquisition circuit (DAS), 210 ... Projection data processing unit, 214 ... Image processing unit

Claims (11)

In an X-ray detector having a plurality of X-ray detection elements arranged along a first direction,
Each of the X-ray detection elements is
A plurality of scintillator portions stacked along a second direction perpendicular to the first direction and having different wavelengths indicating sensitivity peaks with respect to energy of incident X-rays;
An X-ray detector comprising: a plurality of light receiving portions arranged along the first direction on the back surface of the scintillator portion and having different wavelengths indicating sensitivity peaks with respect to incident light. An X-ray generator for generating X-rays;
The X-ray detector that detects X-rays generated from the X-ray generation unit and transmitted through the subject, and has a plurality of X-ray detection elements arranged along a first direction;
A reconstruction processing unit configured to reconstruct at least one image based on an output from the X-ray detector;
Each of the X-ray detection elements is
A plurality of scintillator portions stacked along a second direction perpendicular to the first direction and having different wavelengths indicating sensitivity peaks with respect to energy of incident X-rays;
An X-ray computed tomography apparatus comprising: a plurality of light receiving portions arranged on the back surface of the scintillator portion along the first direction and having different wavelengths indicating sensitivity peaks with respect to incident light. An X-ray generator for generating X-rays;
The X-ray detector that detects X-rays generated from the X-ray generation unit and transmitted through the subject, and has a plurality of X-ray detection elements arranged along a first direction;
A reconstruction processing unit configured to reconstruct at least one image based on an output from the X-ray detector;
Each of the X-ray detection elements is
A plurality of scintillator portions stacked along a second direction perpendicular to the first direction and having different wavelengths indicating sensitivity peaks with respect to energy of incident X-rays;
A plurality of light receiving portions arranged along the first direction on the back surface of the scintillator portion and having different wavelengths indicating sensitivity peaks with respect to incident light;
Each of the light receiving parts is composed of a filter and a photodiode,
An X-ray computed tomography apparatus characterized in that a scintillator portion having a sensitivity peak at the highest X-ray energy among the plurality of scintillator portions is disposed immediately before the light receiving portion.   2. The X-ray detector according to claim 1, wherein a scintillator portion having a sensitivity peak at the highest X-ray energy among the plurality of scintillator portions is disposed immediately before the light receiving portion. The X-ray detector according to claim 1, wherein a scintillator portion having the highest transparency among the plurality of scintillator portions is disposed immediately before the light receiving portion.   The X-ray detector according to claim 1, wherein the light receiving portion is made of a photodiode.     The X-ray detector according to claim 1, wherein each of the light receiving portions includes a filter and a photodiode. Each of the light receiving parts is composed of a filter and a photodiode,
3. The X-ray computed tomography apparatus according to claim 2, further comprising: an X-ray detector in which a scintillator portion having the highest transparency among the plurality of scintillator portions is disposed immediately before the light receiving portion.   The reconstruction processing unit reconstructs a plurality of images related to a wavelength indicating a sensitivity peak of the scintillator portion with respect to the energy of the incident X-ray based on an output from the X-ray detector. The X-ray computed tomography apparatus according to claim 2. Each of the X-ray detection elements has a first light receiving portion and a second light receiving portion,
The reconstruction processing unit reconstructs a first image based on outputs from the plurality of first light receiving portions in the X-ray detector, and a plurality of the second light receiving portions in the X-ray detector. The X-ray computed tomography apparatus according to claim 2, wherein the second image is reconstructed based on an output from the computer. Each of the X-ray detection elements has a first light receiving portion and a second light receiving portion, and the output from the first light receiving portion and the output from the second light receiving portion are integrated as one output. And
The X-ray computed tomography apparatus according to claim 2, wherein the reconstruction processing unit reconstructs one image based on an output from each of the X-ray detection elements.

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