JP2010220880A - Radiation detector and x-ray ct apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector and an X-ray CT apparatus which achieve a higher productivity while restricting the deterioration in detection data. <P>SOLUTION: The radiation detector includes a collimator, a scintillator which is installed facing the collimator and emits fluorescence receiving radiation and photoelectric conversion means which have photoelectric convertors for converting the fluorescence to electric signals and are installed on the main surface of the scintillator. The collimator has the first area where one set of collimator plates is arrayed for one photoelectric convertor and the second area where one set of collimator plates for a plurality of photoelectric converters is arrayed. <P>COPYRIGHT: (C)2011,JPO&INPIT

Description

  The present invention relates to a radiation detector and an X-ray CT apparatus.

In recent X-ray CT (Computer Tomography) apparatuses, a solid detector (hereinafter referred to as a radiation detector) using a scintillator is used in order to increase the number of detection points and increase the spatial resolution.
The radiation detector includes a plurality of photoelectric conversion elements provided on a substrate and a scintillator stacked thereon, and the scintillator is separated by a light reflecting portion for each detection section of the photoelectric conversion elements. Is partitioned. In addition, a collimator is provided to control X-rays incident on individual scintillators and absorb scattered X-rays to reduce crosstalk due to scattered X-rays.

  The collimator plate constituting the collimator is generally formed from a heavy metal such as W (tungsten) or Mo (molybdenum). For this reason, processing is very difficult, which hinders productivity improvement and price reduction of collimators and, consequently, radiation detectors. In particular, in recent years, in order to increase the number of detection points and increase the spatial resolution, the number of rows is increased, and the number of collimator plates arranged tends to increase. For this reason, improving productivity and price reduction of collimators have become important issues.

Here, a technique for arranging a plurality of photoelectric conversion elements with respect to a set of collimator plates has been proposed (see Patent Document 1).
According to the technique disclosed in Patent Document 1, since the number of collimator plates can be reduced, it is possible to improve the productivity and reduce the price of the collimator and, consequently, the radiation detector.
However, if a plurality of photoelectric conversion elements are arranged on a set of collimator plates, X-rays are likely to enter the adjacent scintillators, and crosstalk is likely to occur. As a result, the accuracy of the detected data may be degraded. Further, if the radiation detector having such a configuration is used in an X-ray CT apparatus, the image quality of the X-ray CT image may be deteriorated.

JP 2004-93489 A

  The present invention provides a radiation detector and an X-ray CT apparatus capable of suppressing deterioration in accuracy of detection data and improving productivity.

  According to one aspect of the present invention, the main surface of the scintillator includes a collimator, a scintillator that is provided facing the collimator and emits fluorescence upon receiving radiation, and a photoelectric conversion element that converts the fluorescence into an electrical signal. The collimator includes a first region in which a pair of collimator plates is disposed for one photoelectric conversion element, and a plurality of the photoelectric conversion elements. And a second region having a set of collimator plates disposed thereon. A radiation detector is provided.

  According to another aspect of the invention, an X-ray source, the radiation detector that detects X-rays emitted from the X-ray source, and the X-ray source and the radiation detector are supported. A rotating ring that rotates around the subject and a reconstruction device that reconstructs a tomographic image of the subject based on the intensity of the X-rays detected by the radiation detector. An X-ray CT apparatus is provided.

  ADVANTAGE OF THE INVENTION According to this invention, the radiation detector and X-ray CT apparatus which can aim at the improvement of productivity while being able to suppress the precision degradation of detection data are provided.

It is a mimetic diagram for illustrating a radiation detector concerning a 1st embodiment. It is a schematic cross section for showing the AA cross section in FIG. It is a schematic cross section for showing the BB cross section in FIG. It is a schematic diagram for illustrating the case where a grid ratio is made substantially the same. FIG. 5 is a schematic diagram for illustrating the state of removing scattered X-rays in the case of FIG. FIG. 5 is a schematic diagram for illustrating the state of removing scattered X-rays in the case of FIG. It is a schematic diagram for illustrating the radiation detector which concerns on 2nd Embodiment. It is a schematic diagram for illustrating the radiation detector which concerns on 3rd Embodiment. It is a model perspective view for illustrating the radiation detector concerning a 4th embodiment. 1 is a schematic block diagram for illustrating a schematic configuration of an X-ray CT apparatus.

Hereinafter, embodiments of the present invention will be illustrated with reference to the drawings. In addition, in each drawing, the same code | symbol is attached | subjected to the same component and detailed description is abbreviate | omitted suitably.
The radiation detector according to the present embodiment can be applied to various types of radiation such as γ-rays in addition to X-rays. As a typical example of radiation, the case of X-rays is taken as an example. explain. Therefore, when applying to other radiation in the following embodiments, “X-ray” may be replaced with “radiation”.

FIG. 1 is a schematic diagram for illustrating the radiation detector according to the first embodiment.
FIG. 2 is a schematic cross-sectional view for illustrating an AA cross section in FIG. 1.
FIG. 3 is a schematic cross-sectional view for illustrating a BB cross section in FIG. 2.
As shown in FIGS. 1 to 3, the radiation detector 1 includes a detection unit 2 and a collimator 10. The detection unit 2 includes a scintillator 4, a light reflection unit 17, an adhesive layer 3, a photoelectric conversion unit 12, a signal processing circuit 18, and a base unit 7. The collimator 10 is provided with collimator plates 5 and 5a and holding means 6 for holding the collimator plates 5 and 5a.
In addition, the arrow in a figure represents the incident direction of X-rays.

  As shown in FIG. 3, the scintillator 4 is partitioned corresponding to the detection section of the photoelectric conversion element 12 a provided in the photoelectric conversion means 12, and a groove 16 is formed between the detection sections. That is, each scintillator 4 is divided by the groove 16. And the scintillator 4 and the photoelectric conversion means 12 are joined via the contact bonding layer 3 so that a mutual division may be matched.

  The scintillator 4 is provided facing the collimator 10 and emits fluorescence upon receiving radiation such as X-rays. The fluorescence is, for example, light such as visible light. The scintillator 4 has a maximum light emission wavelength, an attenuation time, a reflection coefficient, a density, a light output ratio, a temperature dependency of the fluorescence efficiency, and the like depending on the material, so that the material can be selected according to the characteristics of each application. it can. As what is used for X-ray CT (Computer Tomography) apparatus, the ceramic scintillator which consists of a sintered compact of rare earth oxysulfide can be illustrated, for example. However, the present invention is not limited to this, and can be changed as appropriate.

Further, the groove 16 between the scintillators 4 is inserted and bonded with a function of reflecting light having a wavelength in the vicinity of the emission wavelength of the scintillator 4 (for example, the illustrated white plate-like body 17a). The light reflection part 17 which consists of these etc. is provided.
The light reflecting section 17 that partitions the scintillator 4 for each photoelectric conversion element 12a plays a role of suppressing optical crosstalk between the sections by performing optical separation and reflection between the sections of each scintillator 4. Yes. In addition, the light reflection part 17 is not necessarily limited to what consists of what inserted and adhere | attached the white plate-shaped body 17a illustrated. For example, it may be made of a white adhesive, or may be filled and solidified with a white pigment. Moreover, it is not necessarily limited to a white thing, What is necessary is just to have a function which reflects the light of the wavelength vicinity of the light emission wavelength of the scintillator 4. FIG. Moreover, the light reflection part 17 may have a role which joins each scintillator 4 and unites them.

  The photoelectric conversion means 12 includes a photoelectric conversion element 12 a that converts fluorescence from the scintillator 4 into an electric signal, and is provided on the main surface of the scintillator 4. As the photoelectric conversion element 12a provided in the photoelectric conversion means 12, for example, a silicon photodiode having a pin structure can be exemplified. The photoelectric conversion element 12a receives the output light corresponding to the section of the scintillator 4 and converts it into an electrical signal. The photoelectric conversion means 12 is not limited to the one provided with the silicon photodiode, and a means for converting the output light from the scintillator 4 into an electric signal (for example, a CCD (Charge Coupled Device)) is appropriately selected. can do.

  The adhesive layer 3 is made of, for example, a transparent adhesive, and is bonded to the scintillator 4 and the photoelectric conversion means 12 while improving light transmission. Thus, each scintillator 4 is joined via the transparent adhesive layer 3 so as to face the light receiving portion of the photoelectric conversion element 12a.

A signal processing circuit 18 is provided on the surface of the photoelectric conversion means 12 that faces the surface to which the scintillator 4 is bonded. The signal processing circuit 18 can capture an electric signal for each section. Further, the signal processing circuit 18 can be provided with an amplifier, an AD converter, etc. (not shown).
Note that the photoelectric conversion means 12 and the signal processing circuit 18 are not necessarily joined, and may be provided separately. Further, an amplifier, an AD converter, and the like (not shown) can be provided separately.

  The base portion 7 has a flat plate shape, and the scintillator 4 provided with the signal processing circuit 18, the photoelectric conversion means 12, the adhesive layer 3, and the light reflecting portion 17 is laminated on the main surface thereof. Moreover, it can attach to the holding means 6 mentioned later using fastening means, such as a screw which is not shown in figure. Therefore, by attaching the base portion 7 to the holding means 6, the scintillator 4 provided so as to be stacked can be attached to the holding means 6.

  The holding means 6 is provided with a groove 6a having a dimension slightly larger than the thickness dimension of the collimator plates 5 and 5a. A plurality of sets of grooves 6a are provided so as to face each other. The collimator plates 5 and 5a can be positioned and held by inserting the end portions of the collimator plates 5 and 5a into the grooves 6a. Further, the groove 6 a is provided so as to have the same pitch dimension as that of the light reflecting portion 17.

  The collimator plates 5 and 5a play a role of controlling X-rays incident on each scintillator 4 and absorbing scattered X-rays to reduce crosstalk caused by the scattered X-rays. The collimator plates 5 and 5a can be made of, for example, W (tungsten), Mo (molybdenum), Ta (tantalum), Pb (lead), or an alloy containing at least one of these heavy metals. However, the material is not limited to these, and a material having excellent X-ray shielding characteristics can be appropriately selected.

  Here, since the collimator plates 5 and 5a are made of heavy metal such as W (tungsten) or Mo (molybdenum), it is very difficult to process. Therefore, productivity is low and production cost is high. This hinders the productivity improvement and price reduction of the collimator 10 and thus the radiation detector 1. In particular, in recent years, in order to increase the number of detection points and increase the spatial resolution, the number of rows is increased, and the number of collimator plates arranged tends to increase. For this reason, improving productivity and price reduction of collimators have become important issues.

In this case, if a plurality of photoelectric conversion elements 12a are arranged on a set of collimator plates, the number of collimator plates can be reduced, so that the production of the collimator 10 and thus the radiation detector 1 can be reduced. Can improve performance and reduce prices.
However, if a plurality of photoelectric conversion elements 12a are simply disposed on a set of collimator plates, the size between the collimator plates increases, and scattered X-rays are likely to enter. Therefore, in the present embodiment, the height dimension of the collimator plate is changed so that the “grid ratio G” is substantially the same even when a plurality of photoelectric conversion elements 12 a are provided.

FIG. 4 is a schematic diagram for illustrating the case where the grid ratios are substantially the same. 4A shows a case where one photoelectric conversion element 12a is provided for a set of collimator plates 5. FIG. 4B shows a case where two photoelectric conversion elements 12a are set for a set of collimator plates 5a. Is provided. In the figure, θ is the allowable incident angle of scattered X-rays, H1 and H2 are the height dimensions of the collimator plates, and W1 and W2 are the dimensions between the collimator plates.
Further, the grid ratio G1 in the case of FIG. 4A is H1 / W1. On the other hand, the grid ratio G2 in the case of FIG. 4B is H2 / W2.

  Here, in order to make the removal effect on the scattered X-rays equal, the allowable incident angle θ of the scattered X-rays may be made substantially the same. In order to make the allowable incident angles θ of scattered X-rays substantially the same, the grid ratios may be made substantially the same, and the height dimension of the collimator plate such that H1 / W1 = H2 / W2. do it.

  For example, since the dimension W2 between the collimator plates is approximately twice the dimension W1, the height dimension H2 of the collimator plates may be approximately twice the dimension H1.

FIG. 5 is a schematic diagram for illustrating how the scattered X-rays are removed in the case of FIG. FIG. 5A is a schematic diagram for illustrating the arrangement of a collimator plate and the like. FIG. 5B is a graph for illustrating the amount of scattered X-rays and direct rays that have reached the detection unit. In the figure, “mountain” represents the amount of the direct line x1, and “valley” represents the amount of the scattered X-ray x2. A line x20 connecting “valleys” represents the distribution of scattered X-rays x2.
FIG. 6 is a schematic diagram for illustrating how the scattered X-rays are removed in the case of FIG. FIG. 6A is a schematic diagram for illustrating the arrangement of a collimator plate and the like. FIG. 6B is a graph for illustrating the amount of scattered X-rays and direct rays that have reached the detection unit. In the figure, “mountain” represents the amount of the direct line x1, and “valley” represents the amount of the scattered X-ray x2. A line x21 connecting “valleys” represents the distribution of scattered X-rays x2.
In this case, the dimensional relationship is such that the grid ratio G1 (H1 / W1) illustrated in FIG. 5 and the grid ratio G2 (H2 / W2) illustrated in FIG. 6 are substantially the same. The water phantom W shown in the figure is an object mainly composed of water, and has a function that approximates the absorption and scattering of X-rays when passing through the human body.

As shown in FIG. 5A, the X-rays exposed from the X-ray tube 101 toward the water phantom W are absorbed and scattered by the water phantom W, and a shielding plate 120 for removing the scattered X-rays. Is incident on. Then, a part of the scattered X-rays is absorbed and removed by the shielding plate 120, and the direct line x1 and the scattered X-ray x2 that has not been removed reach the position where the collimator plate 5 is disposed.
Most of the scattered X-rays x2 are absorbed and removed by entering the collimator plate 5, and a part of the scattered X-rays x2 and the direct line x1 that have not been absorbed reach the detection unit.
Similarly, in the case shown in FIG. 6A, the X-rays emitted from the X-ray tube 101 toward the water phantom W are absorbed and scattered by the water phantom W to remove the scattered X-rays. Is incident on the shielding plate 120. Then, a part of the scattered X-rays is absorbed and removed by the shielding plate 120, and the direct line x1 and the scattered X-ray x2 that has not been removed reach the position where the collimator plate 5a is disposed.
Most of the scattered X-rays x2 are absorbed and removed by entering the collimator plate 5a, and a part of the scattered X-rays x2 and the direct line x1 that have not been absorbed reach the detection unit.

  Here, as shown in FIGS. 5B and 6B, when the grid ratio G1 (H1 / W1) and the grid ratio G2 (H2 / W2) are substantially the same, the detection unit is reached. It can be seen that the distribution state of the scattered X-rays x2 (line x20, line x21) is substantially the same. This means that if the grid ratio is substantially the same, the effect of removing scattered X-rays can be made substantially the same. For this reason, if a plurality of photoelectric conversion elements are disposed on a set of collimator plates with substantially the same grit ratio, the number of collimator plates disposed can be reduced by that amount. It is possible to improve the productivity of the detector and reduce the price.

  However, if a plurality of photoelectric conversion elements are arranged on a set of collimator plates, X-rays are likely to enter the adjacent scintillators, and crosstalk is likely to occur. As a result, the accuracy of the detected data may be degraded. Further, if the radiation detector having such a configuration is used in an X-ray CT apparatus, the image quality of the X-ray CT image may be deteriorated.

As a result of further studies, the present inventor has taken into consideration the positional relationship with the detection target (the subject in the case of the X-ray CT apparatus), and the accuracy of the entire detection data even if the number of collimator plates is reduced. The knowledge that deterioration can be suppressed was obtained.
For example, when a radiation detector is used in an X-ray CT apparatus (see, for example, FIG. 10), there is a portion that is allowed even if the accuracy of detection data is somewhat degraded depending on the positional relationship with the subject. That is, since an important detection site such as the internal organ of the subject exists in the central region of the effective visual field region FOV, it is necessary to increase the accuracy of detection data in the central region. On the other hand, since the possibility that an important detection site exists in the end region of the effective visual field region FOV is low, it is less important than the central region. Therefore, if the number of collimator plates is not reduced in the central area of the effective visual field area FOV and the number of collimator plates is reduced only in the end area of the effective visual field area FOV, The number of collimator plates disposed can be reduced while suppressing deterioration in accuracy.

In the case of the radiation detector 1 illustrated in FIG. 1, a region where a pair of collimator plates 5 is disposed for one photoelectric conversion element 12 a is provided in the central portion of the collimator 10. Further, an area where a pair of collimator plates 5 a is disposed for a plurality of photoelectric conversion elements 12 a is provided in the vicinity of the end of the collimator 10. That is, the number of collimator plates 5a disposed in the end region of the collimator 10 is reduced.
As described above, the collimator 10 includes a region (central region) in which a set of collimator plates 5 is disposed for one photoelectric conversion element 12a and a set of collimator plates 5a for a plurality of photoelectric conversion elements 12a. And a region (end region) in which is disposed. Further, the grit ratio in the center region and the grit ratio in the end region are substantially the same. That is, the center region and the end region are substantially the same with respect to the effect of removing scattered X-rays.

Next, the operation of the radiation detector 1 will be illustrated.
As shown in FIG. 3, X-rays emitted from an X-ray source (not shown) are incident along the collimator plate 5 and reach the scintillator 4 through a space formed between the collimator plates 5. At this time, most of the X-rays (scattered X-rays) incident from a direction different from the direction in which the collimator plate 5 is disposed are absorbed by the collimator plate 5.

Then, the X-rays that have reached the scintillator 4 are converted into light having an intensity proportional to the intensity of the X-rays. The converted light is incident on the photoelectric conversion means 12 while being repeatedly reflected at the surface of the light reflecting portion 17, the interface between the scintillator 4 and the light reflecting portion 17, or the like.
The light incident on the photoelectric conversion means 12 is photoelectrically converted and output as an electric signal having an intensity proportional to the intensity of the light.

  In the present embodiment, the number of collimator plates 5a disposed in the end region of the collimator is reduced. In other words, in the end region of the collimator, a plurality of photoelectric conversion elements 12a are arranged for a set of collimator plates 5a. Further, the grit ratio in the center region and the grit ratio in the end region are substantially the same. That is, the center region and the end region are substantially the same with respect to the effect of removing scattered X-rays.

  Therefore, it is possible to reduce the number of collimator plates disposed while suppressing deterioration in accuracy in the entire detection data. As a result, it is possible to suppress deterioration in accuracy of the detected data and improve productivity and reduce the price.

FIG. 7 is a schematic view for illustrating the radiation detector according to the second embodiment. 7A is a schematic diagram for illustrating the configuration of the radiation detector, FIG. 7B is a schematic enlarged view of a portion C in FIG. 7A, and FIG. 7C is FIG. It is a model enlarged view of the D section in FIG.
As shown in FIG. 7A, the radiation detector 50 includes a detection unit 52 and a collimator 51. Further, as shown in FIGS. 7B and 7C, the detection unit 52 includes the scintillator 4, the light reflection unit 17, the adhesive layer 3, the photoelectric conversion unit 12, and signal processing in the same manner as the detection unit 2 described above. A circuit 18 and a base 7 are provided.

The collimator 51 is provided with a collimator plate 55a to 55c and a holding means (not shown) that holds the collimator plates 55a to 55c.
Also in the present embodiment, the number of collimator plates 55a and 55c disposed in the end region of the collimator is reduced. That is, in the portion C on the most end side, as shown in FIG. 7B, three photoelectric conversion elements 12a are arranged for a set of collimator plates 55a. Further, in the D part adjacent to the C part, as shown in FIG. 7C, two photoelectric conversion elements 12a are arranged for a set of collimator plates 55b. In the central region, one photoelectric conversion element 12a is arranged for one set of collimator plates 55c.
And the grit ratio of the collimator in the center region is substantially the same as the grit ratio of the collimator in the C part and the D part. That is, in the region where a set of collimator plates is disposed for the plurality of photoelectric conversion elements 12a, the height dimension of the collimator plates increases stepwise toward the end of the collimator 51.
In this way, if the grid ratio is substantially the same, the effect of removing scattered X-rays in the central region and the effect of removing scattered X-rays in the C and D portions can be made substantially the same.

Further, by reducing the number of collimator plates 55a in the less important C portion than the number of collimator plates 55b in the D portion disposed on the central region side, it is possible to reduce more collimator plates. I am trying.
In this way, if the number of collimator plates to be reduced is increased stepwise from the central region toward the end region, the accuracy change of the detection data can be made stepwise. Therefore, continuity in the accuracy change of the detection data can be improved, so that deterioration in accuracy in the entire detection data can be further suppressed.
Since the operation of the radiation detector 50 is the same as that of the radiation detector 1 described above, the description thereof is omitted.

  According to the present embodiment, it is possible to reduce more collimator plates. In addition, by increasing the number of collimator plates in a stepwise manner from the center region toward the end region, the accuracy change of the detection data can be made stepwise. Therefore, it is possible to further suppress deterioration in accuracy in the entire detection data.

FIG. 8 is a schematic view for illustrating the radiation detector according to the third embodiment.
As shown in FIG. 8, the radiation detector 60 includes a detection unit 62 and a collimator 61. Similarly to the detection unit 2 described above, the detection unit 62 includes a scintillator, a light reflection unit, an adhesive layer, a photoelectric conversion unit, a signal processing circuit, and a base (not shown).
The collimator 61 is provided with a collimator plate 65 and a holding means (not shown) that holds the collimator plate 65.

Also in the present embodiment, the number of collimator plates 65 disposed in the end region of the collimator is reduced. However, in the present embodiment, the number of collimator plates 65 gradually decreases from the central region toward the end region. In addition, the grit ratio of the collimator in the center region and the grit ratio of the collimator in the region where the number of collimator plates 65 gradually decreases are substantially the same. That is, in the region where a set of collimator plates is disposed for the plurality of photoelectric conversion elements 12 a, the height dimension of the collimator plate 65 gradually increases toward the end of the collimator 61.
In this way, if the number of collimator plates 65 is gradually decreased from the center region toward the end region, the continuity in the accuracy change of the detection data can be further improved. Therefore, it is possible to further suppress deterioration in accuracy in the entire detection data.
Since the operation of the radiation detector 60 is the same as that of the radiation detector 1 described above, the description thereof is omitted.

  According to the present embodiment, it is possible to reduce more collimator plates. Moreover, the accuracy change of detection data can be made more continuous by gradually decreasing the number of collimator plates from the central region toward the end region. Therefore, it is possible to further suppress deterioration in accuracy in the entire detection data.

  FIG. 9 is a schematic perspective view for illustrating a radiation detector according to the fourth embodiment. In the figure, the X direction and the Y direction represent two directions orthogonal to each other. When the radiation detector 70 is provided in the X-ray CT apparatus, the X direction is the channel direction and the Y direction is the slice direction. Moreover, the arrow in a figure represents the incident direction of X-rays.

As shown in FIG. 9, the radiation detector 70 includes a detection unit 72 and a collimator 71. The detector 72 is provided with the scintillator 4, the light reflector 17, the adhesive layer 3, the photoelectric conversion means 12, the signal processing circuit 18, and the base 7. The collimator 71 is provided with a collimator plate 75 and a holding means (not shown) that holds the collimator plate 75.
The scintillator 4 is partitioned corresponding to the detection section of the photoelectric conversion element 12 a provided in the photoelectric conversion means 12, and a light reflecting portion 17 is provided between the scintillators 4. In the present embodiment, it is a lattice-like light reflecting portion 17 that partitions the scintillator 4 for each photoelectric conversion element 12 a provided in the photoelectric conversion means 12.
Further, the collimator 71 is also formed in a lattice shape so as to correspond to the lattice-like section of the scintillator 4. That is, the collimator 71 is provided with a collimator plate 75 in a lattice shape.

Also in the present embodiment, the number of collimator plates 75 disposed in the end region of the collimator is reduced. That is, the end region in the X direction (channel direction when the radiation detector 70 is provided in the X-ray CT apparatus) and the end in the Y direction (slice direction when the radiation detector 70 is provided in the X-ray CT apparatus). In the partial area, a plurality of photoelectric conversion elements 12 a are arranged for a set of collimator plates 75. Further, by having a grit ratio substantially the same as the grit ratio of the collimator in the central region, the effect of removing scattered X-rays is made substantially the same as that in the central region.
Since the operation of the radiation detector 70 is the same as that of the radiation detector 1 described above, the description thereof is omitted.

  According to the present embodiment, it is possible to reduce the number of elements of collimator 71 (the number of collimator plates 75) formed in a lattice while suppressing accuracy deterioration in the entire detection data. As a result, it is possible to suppress deterioration in accuracy of the detected data and improve productivity and reduce the price.

Next, an X-ray CT apparatus according to this embodiment is illustrated.
FIG. 10 is a schematic block diagram for illustrating a schematic configuration of the X-ray CT apparatus.
As shown in FIG. 10, the X-ray CT apparatus 100 includes an imaging unit 100a and a processing / display unit 100b.

  The imaging unit 100a exposes the subject to X-rays, detects the X-rays that have passed through the subject, and obtains projection data (or raw data). The imaging means is a rotation / rotation (ROTATE / ROTATE) type in which the X-ray tube and the detector unit rotate as a unit, and a plurality of detection elements are attached in a ring shape, and only the X-ray tube There are various types, such as a fixed / rotated (STATIONARY / ROTATE) type that rotates around the subject, and a type that electronically moves the position of the X-ray source on the target by deflecting the electron beam. Even if it is a type, the radiation detector according to the present embodiment can be applied. Here, as an example, a rotation / rotation type X-ray CT apparatus will be described as an example.

  As shown in FIG. 10, the imaging means 100a includes an X-ray tube 101, a rotating ring 102, a detector unit 103, a data acquisition circuit (DAS) 104, a non-contact data transmission device 105, a gantry driving unit 107, and a slip ring 108. The radiation detector 1 is provided.

An X-ray tube 101 that is an X-ray source is a vacuum tube that generates X-rays, and is supported by a rotating ring 102. The X-ray tube 101 is supplied with power (tube current, tube voltage) necessary for X-ray exposure from the high voltage generator 109 via the slip ring 108. The X-ray tube 101 emits X-rays toward a subject in the effective visual field region FOV by causing electrons accelerated by the supplied high voltage to collide with the target.
An X-ray (not shown) that shapes the shape of the X-ray beam exposed from the X-ray tube 101 into a cone shape (quadrangular pyramid shape) or a fan beam shape between the X-ray tube 101 and the subject. A tube side collimator is provided.

  The detector unit 103 is a detector system that detects X-rays transmitted through the subject, and is provided on the rotating ring 102 so as to face the X-ray tube 101. The radiation detector 1 is attached to the outer peripheral side of the detector unit 103 (the side opposite to the subject). That is, on the outer peripheral side of the detector unit 103, a holding unit 6 holding the collimator plates 5 and 5a, and a base unit 7 provided with the scintillator 4, the photoelectric conversion unit 12 and the like are attached. Further, as described above, the number of collimator plates 5a is not reduced in the central region of the effective visual field region FOV, but the number of collimator plates 5a is reduced only in the end region of the effective visual field region FOV. ing. In addition, the grit ratio of the collimator in the central region and the grit ratio of the collimator in the end region are made substantially the same.

The X-ray tube 101 and the detector unit 103 are provided on the rotating ring 102. The rotating ring 102 is driven by the gantry driving unit 107 and rotates around the subject.
The data collection circuit (DAS) 104 has a plurality of data collection element arrays in which DAS chips are arranged, and receives data detected by the detector unit 103 (hereinafter referred to as raw data). The input raw data is subjected to amplification processing, A / D conversion processing, and the like, and then transmitted to the preprocessing device 106 provided in the processing / display unit 100b via the data transmission device 105.
The gantry driving unit 107 performs driving such as rotating the X-ray tube 101 and the detector unit 103 integrally around a central axis parallel to the body axis direction of the subject inserted into the diagnostic aperture. Control it.

  Next, the processing / display unit 100b is illustrated. The processing / display unit 100b includes a preprocessing device 106, a high voltage generation device 109, a host controller 110, a storage device 111, a reconstruction device 114, an input device 115, a display device 116, an image processing unit 118, a network communication device 119, data. / A control bus 300 is provided.

  The preprocessing device 106 receives raw data from the data acquisition circuit (DAS) 104 via the data transmission device 105, and executes sensitivity correction and X-ray intensity correction. Note that the raw data preprocessed by the preprocessing device 106 is referred to as “projection data”.

The high voltage generator 109 supplies power necessary for X-ray exposure to the X-ray tube 101 via the slip ring 108. The high voltage generator 109 includes a high voltage transformer, a filament heating converter, a rectifier, a high voltage switch, and the like.
The host controller 110 performs overall control related to various processing such as photographing processing, data processing, and image processing.

The storage device 111 stores image data such as collected raw data, projection data, and CT image data.
The reconstruction device 114 performs reconstruction processing on projection data based on predetermined reconstruction parameters (reconstruction area size, reconstruction matrix size, threshold for extracting a region of interest, etc.), and thereby a predetermined slice. Create reconstructed image data. In general, reconstruction processing includes cone beam reconstruction (Feldkamp method, ASSR method, etc.) and fan beam reconstruction, and any method can be used.
The input device 115 is provided with a keyboard, various switches, a mouse, and the like so that various scanning conditions such as a slice thickness and the number of slices can be input by an operator.

  The image processing unit 118 performs image processing for display, such as window conversion and RGB processing, on the reconstructed image data created by the reconstructing device 114 and outputs the image processing to the display device 116. Further, the image processing unit 118 creates a so-called pseudo three-dimensional image such as a tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, or a three-dimensional surface image based on an instruction from the operator, and outputs the generated image to the display device 116. . The output image data is displayed on the display device 116 as an X-ray CT image.

The network communication device 119 transmits and receives various data to and from other devices and network systems such as RIS (Ragiology Information System) via the network.
The data / control bus 300 is a signal line for connecting various devices and transmitting / receiving various data, control signals, address information, and the like.

Next, the operation of the X-ray CT apparatus 100 according to the present embodiment will be illustrated.
In order to obtain a desired image by photographing the subject inserted into the diagnostic aperture, first, various scanning conditions such as the slice thickness and the number of slices are input from the input device 115.
As the operation of the X-ray CT apparatus 100 starts, the rotating ring 102 starts rotating, and at the same time, X-rays are exposed from the X-ray tube 101 toward the subject.
The X-ray transmitted through the subject reaches the radiation detector 1 of the detector unit 103 provided so as to face the X-ray tube 101 across the subject.
The radiation detector 1 is provided with collimator plates 5 and 5 a to remove scattered X-rays incident from other than the focal direction of the X-ray tube 101. For this reason, light based on X-rays from the focal direction of the X-ray tube 101 is incident on the photoelectric conversion means 12 (photoelectric conversion element 12 a) of the radiation detector 1.

  The light received by the photoelectric conversion means 12 (photoelectric conversion element 12a) is converted into an electrical signal proportional to the intensity and output to the data collection circuit (DAS) 104. The electrical signal (raw data) input to the data acquisition circuit (DAS) 104 is transmitted to the preprocessing device 106 after being subjected to amplification processing, A / D conversion processing, and the like. In the pre-processing device 106, sensitivity correction and X-ray intensity correction are performed on the transmitted raw data to generate projection data. In the reconstruction device 114, reconstructed image data for a predetermined slice is created from projection data based on a predetermined reconstruction parameter. In the image processing unit 118, image processing for display such as window conversion of reconstructed image data and RGB processing is performed and output to the display device 116. Thereby, a tomographic image (slice image) of the subject is obtained. The image processing unit 118 also creates a so-called pseudo three-dimensional image such as a tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, a three-dimensional surface image, etc. based on a command from the operator. Note that raw data, projection data, image data, and the like are stored in the storage device 111.

According to the present embodiment, the number of collimator plates 5a disposed in the collimator in the end region of the effective visual field region FOV is reduced. That is, in the collimator in the end region, a plurality of photoelectric conversion elements 12a are arranged for one set of collimator plates 5a. Further, by having a grit ratio substantially the same as the grit ratio of the collimator in the central region, the effect of removing scattered X-rays is made substantially the same as that in the central region.
Therefore, it is possible to reduce the number of collimator plates disposed while suppressing deterioration in accuracy in the entire detection data. As a result, it is possible to suppress deterioration in accuracy of detection data, and it is possible to improve the productivity and reduce the price of the X-ray CT apparatus 100.

The embodiment of the present invention has been illustrated above. However, the present invention is not limited to these descriptions.
As long as the features of the present invention are provided, those skilled in the art appropriately modified the design of the above-described embodiments are also included in the scope of the present invention.
For example, the shape, size, material, arrangement, number, and the like of each element included in the radiation detector 1, the radiation detector 50, the radiation detector 60, the radiation detector 70, and the X-ray CT apparatus 100 are limited to those illustrated. However, it can be changed as appropriate.
Moreover, although the position which reduces the arrangement | positioning number of a collimator board is made into the edge part area | region, it is not necessarily limited to this. Even if the accuracy of the detection data is somewhat deteriorated due to the content or positional relationship of the detection target, it is possible to reduce the number of collimator plates that are allowed to be provided.
Moreover, each element with which each embodiment mentioned above is combined can be combined as much as possible, and what combined these is also included in the scope of the present invention as long as the characteristics of the present invention are included.

  DESCRIPTION OF SYMBOLS 1 Radiation detector, 2 detection part, 3 adhesion layer, 4 scintillator, 5 collimator board, 5a collimator board, 6 holding means, 7 base, 10 collimator, 12 photoelectric conversion means, 12a photoelectric conversion element, 17 light reflection part, 18 Signal processing circuit, 50 radiation detector, 51 collimator, 52 detector, 55a to 55c collimator plate, 60 radiation detector, 61 collimator, 62 detector, 65 collimator plate, 70 radiation detector, 71 collimator, 72 detector, 75 collimator plate, 100 X-ray CT apparatus, 100a imaging means, 100b processing / display means, 101 X-ray tube, 103 detector section

Claims (7)

A collimator,
A scintillator that is provided facing the collimator and emits fluorescence upon receiving radiation;
A photoelectric conversion element that converts the fluorescence into an electrical signal, photoelectric conversion means provided on the main surface of the scintillator,
With
The collimator includes a first region in which a set of collimator plates is disposed for one photoelectric conversion element, and a second region in which a set of collimator plates is disposed for a plurality of the photoelectric conversion elements. A radiation detector characterized by comprising:   The radiation detector according to claim 1, wherein the grid ratio in the first region and the grid ratio in the second region are substantially the same. The first region is provided in a central portion of the collimator;
The radiation detector according to claim 1, wherein the second region is provided in the vicinity of an end of the collimator.   4. The radiation according to claim 1, wherein a height dimension of the collimator plate is increased in a stepwise manner toward an end of the collimator in the second region. Detector.   The radiation detector according to any one of claims 1 to 3, wherein a height dimension of the collimator plate gradually increases toward the end of the collimator in the second region.   The radiation detector according to any one of claims 1 to 5, wherein the collimator is provided with the collimator plate in a lattice shape. An X-ray source;
The radiation detector according to any one of claims 1 to 6, which detects X-rays emitted from the X-ray source;
A rotating ring that supports the X-ray source and the radiation detector and rotates around a subject;
A reconstruction device that reconstructs a tomographic image of the subject based on the intensity of X-rays detected by the radiation detector;
An X-ray CT apparatus comprising:

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