JPWO2015098631A1 - Radiation detector and X-ray CT apparatus - Google Patents

Abstract

In order to provide a radiation detector having a structure that can easily and accurately shield X-rays directly incident on the photodiode element array and prevent vibration caused by high-speed rotation, the radiation incident surface of the scintillator element array 13 is provided. A shielding portion 17 that shields the reflecting material 12 between the adjacent scintillator elements 11 is provided. The width Ws of the shielding part 17 is smaller than the width Wr of the reflector 12, and larger than the thickness Wc of the collimator plate 16. Further, an adhesive layer 18 is disposed between the upper surface of the shielding part 17 and the end surface of the collimator plate 16, and the upper surface of the shielding part 17 and the end surface of the collimator plate 16 are adhered by an adhesive layer.

Description

  The present invention relates to a radiation detector that detects X-rays, γ-rays, and the like, and more particularly to a medical image diagnostic apparatus such as a CT apparatus using the radiation detector.

  At present, as a radiation detector of an X-ray CT apparatus, an indirect conversion type detector in which a phosphor element such as a ceramic scintillator and a photodiode element are combined is mainly used. This indirect conversion type detector has a detector element module in which a plurality of scintillator elements are arranged in a two-dimensional array, and a photodiode element is arranged on the back of each scintillator element. Many of the structures that are arranged on top are used. The photodiode outputs a current signal corresponding to the X-ray dose that has passed through the subject. This output current signal is converted into a digital signal by the AD converter circuit board, and then transmitted to the image processing device to create a CT image. Is done.

  A plurality of scintillator elements arranged in a two-dimensional array of detection element modules are separated by a reflective material such as a white resin between the scintillator elements, and one scintillator element corresponds to one pixel of a CT image. For this reason, the arrangement pitch of the scintillator elements dominates the spatial resolution of the radiation detector. Further, it is known that the thickness of the reflecting material separating the scintillator elements affects the amount of optical crosstalk between the scintillator elements, and consequently affects the spatial resolution.

  In addition, in order to prevent only scattered X-rays scattered by the subject from being incident on the detection element by direct incidence of the direct X-ray directed from the focal point of the X-ray tube linearly toward the detection element module to the scintillator element. A removal collimator is disposed in front of the detection element module. The scattered radiation removing collimator includes a plurality of collimator plates arranged so that the main plane is substantially perpendicular to the upper surface of the scintillator element. The collimator plates are arranged side by side so as to cover the reflective material in the gap between adjacent scintillator elements. Therefore, the larger the thickness of the collimator plate, the higher the scattered radiation removal capability, while the direct X-rays themselves that should be incident on the detector are also shielded.

Patent Documents 1 and 2 disclose a configuration in which the upper surface of the reflective material between the scintillator elements and the upper surface of the edge of the scintillator element are covered with a shielding member having a width wider than that of the reflective material.
The collimator plate is disposed on the shielding member with a gap from the shielding member. By arranging the shielding member in this way, the reflection material is prevented from deteriorating due to the incidence of X-rays on the reflection material, and the X-ray passes through the reflection material and is detected by the photodiode to generate noise. In addition, the edge of the reflector is prevented from being damaged during the manufacturing process.

  In Patent Document 3, corresponding to the movement of the focal position of the X-ray, in order to remove the scattered radiation incident on each scintillator element in the same way, the collimator plate can be a wedge-shaped cross-sectional shape, It has been proposed to have a laminated structure of plates with different thicknesses.

US Pat. No. 6,934,354 U.S. Pat.No.7010083 Japanese Patent Laid-Open No. 11-218578

  In the structures disclosed in Patent Documents 1 and 2, the upper surface of the reflective material between the scintillator elements and the upper surface of the edge of the scintillator element are covered with a shielding member having a width wider than that of the reflective material. Not only can X-rays to be reduced be reduced, but also X-rays that pass through the scintillator element and enter the reflector from the side surface of the scintillator element can be reduced. However, since the edge of the scintillator element is covered with the shielding member, there is a problem that the aperture ratio of the scintillator element is substantially reduced and the amount of X-rays incident on the scintillator element is reduced. Further, in the technologies disclosed in Patent Documents 1 and 2, the collimator plate and the shielding member are separated, and when used in a high-speed CT apparatus of about 0.3 seconds / revolution in recent years, the collimator plate vibrates. There is a possibility that. Patent Document 1 describes that the collimator plate may be connected to the shielding member, but there is no disclosure of a specific structure, and how the minute shielding member is connected to the collimator plate. Is not disclosed.

  On the other hand, as described in Patent Document 3, the collimator plate has a wedge-shaped cross-sectional shape and the like, and the structure corresponding to the focal movement of the X-rays needs to process the collimator plate into a complicated shape with high accuracy. Incurs an increase. Further, since the thickness of the collimator plate shown in Patent Document 3 is larger than the width of the reflector between the scintillator elements, the edge of the scintillator element is also covered with the collimator plate, and the aperture ratio of the scintillator element is increased. Substantially lower.

  An object of the present invention is to provide a radiation detector having a structure that can easily and accurately shield X-rays directly incident on a photodiode element array and prevent vibration due to high-speed rotation.

  The radiation detector of the present invention includes a shielding portion that shields a reflective material between adjacent scintillator elements on the radiation incident surface of the scintillator element array. The width Ws of the shielding part is smaller than the width Wr of the reflector and larger than the thickness Wc of the collimator plate. An adhesive layer is disposed between the upper surface of the shielding part and the end face of the collimator plate, and the upper surface of the shielding part and the end face of the collimator plate are adhered by the adhesive layer.

  According to the present invention, X-rays directly incident on the photodiode element array can be easily and accurately shielded, and vibration due to high-speed rotation can be prevented.

1 is a cross-sectional view of a detector module according to a first embodiment of the present invention. FIG. 2 is a block diagram showing a configuration in which a detector is provided in the detector module according to the first embodiment. (a) A perspective view of the detector module of FIG. 1, (b) a perspective view of the radiation detector of the first embodiment. The graph which shows the relationship between the dynamic range of the output of the photodiode element of the radiation detector for CT apparatuses of this invention, and the noise component current contained in an output signal current. (a) Explanatory diagram showing direct X-rays incident on the scintillator element 11 of the detector module of the first embodiment, (b) Explanatory diagram showing direct X-rays incident on the scintillator element 11 of the detector module of the comparative example . 3 is a graph showing an example of the output of the photodiode element 14 with respect to X-ray energy in the case of Ws <Wr and Ws> Wr, respectively. 6 is a graph showing fluctuations in the output of the photodiode element due to X-ray components incident on the edge portion of the scintillator element. 6 is a flowchart showing a procedure of energy characteristic correction according to the first embodiment. 6 is a graph showing the relationship between the object size and the output signal of the photodiode element in the detector module of the first embodiment. Sectional drawing of the detector module of 2nd Embodiment. FIG. 11 is a partial cross-sectional view when the detector modules of FIG. 10 are arranged side by side. Sectional drawing of the detector module of 3rd Embodiment. Sectional drawing of the detector module of 4th Embodiment. The block diagram of the X-ray CT apparatus of 5th Embodiment.

  Embodiments of the present invention will be described with reference to the drawings.

(First embodiment)
As shown in FIG. 1, the radiation detector of the present invention comprises a scintillator element array 13, a photodiode element 14, and a collimator plate 16.

  The scintillator element array 13 includes a plurality of scintillator elements 11 that emit fluorescence by radiation and a reflecting material 12, and the scintillator elements 11 are arranged one-dimensionally or two-dimensionally. The reflective material 12 is disposed at least in a region between adjacent scintillator elements 11 and reflects fluorescence emitted inside the scintillator elements 11.

  The photodiode elements 14 are respectively disposed on the surfaces of the plurality of scintillator elements 11 opposite to the radiation incident surface. That is, the photodiode elements 14 constitute a photodiode element array 15 arranged in an array at the same pitch as the scintillator element array 13. The photodiode element 14 detects the fluorescence emitted by the scintillator element 11.

  The collimator plate 16 is disposed above the adjacent scintillator elements 11 on the radiation incident surface side of the scintillator element array 13. The collimator plate 16 absorbs radiation scattered by the subject and prevents the scattered radiation from reaching the scintillator element 11.

  Further, on the radiation incident surface of the scintillator element array 13, a shielding portion 17 that shields the reflecting material 12 between the adjacent scintillator elements 11 is disposed. The width Ws of the shielding part 17 is set to be smaller than the width Wr of the reflector 12 and larger than the thickness Wc of the collimator plate 16. The shielding unit 17 transmits radiation that passes through the subject from the focal point of the radiation source and directly reaches the radiation detector (hereinafter referred to as direct radiation) through the reflector 12 and reaches the photodiode element 14. It is preventing. In addition, by setting the width Ws of the shielding part 17 to be smaller than the width Wr of the reflector 12 and larger than the thickness Wc of the collimator plate 16, the edge of the scintillator element 11 is prevented from being covered by the shielding part 17. The aperture ratio of the radiation detector is prevented from decreasing.

  An adhesive layer 18 is disposed between the upper surface of the shielding part 17 and the end surface of the collimator plate 16. The upper surface of the shielding part 17 and the end surface of the collimator plate 16 are bonded by the adhesive layer 18. Thereby, even when the radiation detector is used in the CT apparatus and rotated at high speed, the collimator plate 16 can be prevented from vibrating, so that it is possible to prevent the occurrence of radiation detection errors due to fluctuations in the position of the collimator 16. .

  The radiation detector of the present invention preferably further includes a correction unit 20 that corrects the output of the photodiode element 14 as shown in FIG. The correcting unit 20 outputs the photodiode element 14 for each of the plurality of scintillator elements 11 due to the radiation that has passed between the collimator plates 16 entering the edge of the scintillator element 11 and being converted into fluorescence. Correct variations in the signal. For example, the correction unit 20 can correct the output signal of the photodiode element 14 for each scintillator element 11 using a correction coefficient predetermined for each scintillator element 11 in order to correct variation. It is desirable that a correction coefficient is prepared for each radiation energy.

Hereinafter, specific examples of the radiation detector according to the first embodiment of the present invention will be described in detail. In the following embodiments, a case will be described in which the radiation detector is a radiation detector of an X-ray CT apparatus. FIG. 1 shows a cross-sectional view of a detector module 101 of the radiation detector of the present invention.
The detector module 101 includes a detection element module 102 and a scattered radiation removal collimator 103.

  The detection element module 102 includes a scintillator element array 13 and a photodiode element array 15. The scintillator element array 13 has a configuration in which a plurality of scintillator elements 11 are arranged in a two-dimensional array with a reflector 12 interposed therebetween. The scintillator element 11 is formed of a phosphor that emits fluorescence when X-rays enter. The reflective material 12 is formed of a material that reflects fluorescence. In the example of FIG. 1, the reflecting material 12 is disposed not only in the gap between adjacent scintillator elements, but also on the upper surface of the scintillator element 11 (the surface on which X-rays are incident), and the fluorescence emitted upward is downward ( Reflected on the photodiode element 14 side).

  Similar to the arrangement of the scintillator element array 13, the photodiode elements 14 are arranged in a two-dimensional array to constitute a photodiode element array 15. The photodiode element array 15 is bonded and fixed to the scintillator element array 13 with a transparent adhesive.

  As shown in FIG. 1, the scattered radiation elimination collimator 103 includes a plurality of collimator plates 16 respectively installed at the positions of the gaps of the scintillator elements 11, and a shielding part 17 disposed on the reflection part 12 of the gaps of the scintillator elements 11. Have The plurality of collimator plates 16 are supported by collimator support portions 31 arranged on both sides of the scintillator element array 13 as shown in the perspective view of FIG. The collimator support unit 31 is supported by the substrate 115. The direction of the collimator plate 16 is determined so that the focal point of the X-ray source is positioned on the extended line of the main plane.

  The lower end of the collimator plate 16 is positioned and supported by a collimator support 31 so that a predetermined gap is generated with respect to the upper surface of the shield 17. The adhesive layer 18 disposed so as to wrap the shielding part 17 adheres the lower end of the collimator plate 16 and the upper surface of the shielding part 17 and also adheres and fixes the shielding part 17 and the scintillator element array 13.

  The shielding part 17 may be formed by printing a resin mixed with tungsten or molybdenum powder on the scintillator element array 13. Alternatively, the shielding part 17 may be formed by bonding the resin molded into a predetermined shape onto the scintillator element array 13. You may do it. Alternatively, the shielding portion 17 may be formed by bonding a metal sheet of tungsten or molybdenum to a predetermined size and bonding it onto the scintillator element array 13. As the collimator plate 16, for example, a heavy metal plate such as tungsten or molybdenum is used.

  The radiation detector of the X-ray CT apparatus has a structure in which a plurality of detector modules 101 are arranged in the x direction and supported by polygons 120 as shown in FIG. Thus, the detector modules 101 are arranged in an arc shape so as to face the X-ray tube focus.

  Next, the operation of the detector module 101 of the present invention will be described. Direct X-rays irradiated from the focal point of the X-ray tube (not shown in FIG. 1) and transmitted through the subject (not shown) enter the detector module 101 along the z direction from above in FIG. It passes between 16 and is absorbed by the scintillator element 11. At this time, the scattered X-rays scattered by the subject are incident on the detector module 101 at a certain angle with respect to the z direction, so that they are absorbed by the collimator plate 16 and do not reach the scintillator element 11. The X-rays absorbed by the scintillator element 11 are converted into visible light fluorescence and detected by the photodiode element 14. The photodiode element 14 generates an analog electric signal corresponding to the emission intensity.

  An analog electric signal generated by the photodiode element 14 is converted into a digital signal through an AD conversion circuit in a detector circuit 316 disposed on the lower surface of the substrate 115, and is corrected by the correction unit 20 in FIG. The corrected digital signal is used for image processing for reconstructing a CT image.

  At this time, in the present embodiment, the direct X-rays that have passed through the collimator plate 16 are reflected by the reflector 12 by installing a shielding portion 17 having a width Ws that is larger than the width Wc of the collimator plate 16 and smaller than the width Wr of the reflector 12. And is incident on the photodiode element array 15. The reason why the width Ws of the shielding part is made smaller than the width Wr of the reflector 12 is that the shielding part 17 does not shield the edge part of the scintillator element 11. Thereby, the shielding part 17 does not reduce the amount of direct X-rays that should be absorbed by the scintillator element 11, and it is possible to prevent the aperture ratio as a radiation detector from being lowered. Furthermore, by making the width Ws of the shielding part 17 smaller than the width Wr of the reflecting material 12, the shielding part 17 covers the edge (edge part) of the scintillator element 11 due to the dimensional tolerance of each part and the assembly tolerance. It is possible to reduce the possibility of being disposed or being disposed at a position that is not covered, and to suppress characteristic variation between pixels.

  Here, X-ray shielding by the shielding unit 17 and variation correction by the correction unit 20 will be further described.

  FIG. 4 is a graph showing the relationship between the dynamic range of the output of the photodiode element 14 of the radiation detector for CT apparatus and the noise component current included in the output signal current, and the cause of noise. As is clear from FIG. 4, in the region where the output signal current is extremely weak, the circuit noise is dominant, but in the region of the output signal current that is larger than a certain level, the quantum noise is dominant, and the CT image It can be seen that most of the noise is quantum noise. Therefore, in the present invention, the shielding unit 17 is designed to reduce quantum noise.

  FIG. 5A shows direct X-rays incident on the scintillator element 11 when the width Ws of the shielding portion 17 of the present embodiment is Ws <Wr with respect to the width Wr of the reflector 12. FIG. 5B shows direct X-rays incident on the scintillator element 11 when Ws> Wr as a comparative example. As shown in FIG. 5A, when Ws <Wr, direct X-rays in the angle range from B0 to B2 enter the scintillator element 11. In particular, a component between B1 and B2, which is a part of direct X-rays that should be incident on the scintillator element 11, can be incident on the scintillator element 11.

  In contrast, as in the comparative example of FIG. 5B, when Ws> Wr, the component between B1 and B2 is shielded by the shielding part 17 and cannot be incident on the scintillator element 11. .

  FIG. 6 is a graph showing an example of the output of the photodiode element 14 with respect to incident X-ray energy in the case of Ws <Wr and Ws> Wr. In these examples, Wr is constant in order to keep the optical crosstalk amount constant.

  As is apparent from the graph of FIG. 6, when Ws> Wr of the comparative example, a part of preferable direct X-rays that should be incident on the scintillator element 11 is shielded, so that the output of the photodiode element 14 decreases. It is done. This decrease in output means a decrease in the aperture ratio of the X-ray detector and causes a decrease in the S / N ratio accompanying an increase in quantum noise. As a result, the quality of the reconstructed CT image is degraded. On the other hand, when Ws <Wr of the present invention, since the aperture ratio increases, the output of the photodiode element 14 also increases, the quantum noise decreases, and a good CT image with a high SN ratio is obtained. I can do it.

  However, as is apparent from FIG. 6, the change in the output of the photodiode element 14 described above exhibits non-linearity because the X-ray energy dependence characteristic is not linear. This is because direct X-rays (components between B1 and B2) that are not shielded by setting Ws <Wr in the present invention are components that enter the edge (edge) of the scintillator element 11. is there. Therefore, the relationship between the nonlinearity of the output of the photodiode element 14 and the shape of the edge portion (edge portion) of the scintillator element 11 was examined in more detail, and the graph of FIG. 7 was obtained.

  FIG. 7 is a graph showing fluctuations in the output of the photodiode element 14 due to the X-ray component incident on the edge portion of the scintillator element 11. FIG. 7 shows the ratio between the output when the X-ray is irradiated to the edge (edge) of the scintillator element 11 and the output when the X-ray is not irradiated to the edge (edge) of the scintillator element 11. Show. That is, the output fluctuation of the photodiode element 14 when the edge portion becomes the smallest due to variations in processing accuracy or the like is shown for each X-ray energy. When the edge part shape of the scintillator element 11 fluctuates, the absorbed dose of X-rays absorbed at the edge part fluctuates, but the rate at which the absorbed dose fluctuates is not constant depending on the X-ray energy. As a result, it can be seen that the ratio of the output fluctuation to each X-ray energy is not constant as shown in FIG. Further, as compared with the scintillator element at the center of the detection element module, the element at the end of the module increases the proportion of X-rays incident on the edge portion, and thus the influence of characteristic variation becomes significant.

  This nonlinear output fluctuation with respect to the X-ray energy varies for each photodiode element 14 and may appear as an artifact in the CT image. For example, the energy of transmitted X-rays that have passed through the subject is surfaced by a phenomenon called beam hardening that varies depending on the size of the subject. That is, the average energy of transmitted X-rays shifts to the higher energy side as the subject is larger. Therefore, in order to accurately measure the X-ray absorption coefficient of the subject, it is necessary to correct the output fluctuation generated in each detection element for each individual photodiode element. Since this output fluctuation correction needs to correct nonlinear fluctuations as shown in FIGS. 6 and 7, it cannot be corrected by conventional linearity sensitivity correction such as fluctuation of aperture ratio.

  Therefore, in the present embodiment, nonlinear fluctuations in the output signal are corrected for each photodiode element 14 by the correction unit 20 arranged in the detector circuit 316. Specifically, output correction is performed on the output of the photodiode element 14 in consideration of the amount of energy shift due to transmission through the subject.

  The correction procedure will be described more specifically with reference to FIG. FIG. 8 is a flowchart showing the procedure of energy characteristic correction. First, at the time of product shipment or regular inspection, a plurality of phantoms simulating subjects of various sizes are used, and these are irradiated with X-rays and detected by a radiation detector, whereby a plurality of photodiode elements 14 are obtained. Are respectively measured (steps 801 and 802). As a result, as shown in FIG. 9, the relationship between the object size and the output signal is obtained for each photodiode element. Based on the relationship between the obtained object size and the output signal, a correction coefficient for output correction taking into account the amount of energy shift due to transmission through the object is calculated based on the correction equation. This correction coefficient is obtained for each radiation energy, and a correction coefficient table representing the relationship between the output of the photodiode element 14 and the correction coefficient is created (step 803). This correction coefficient table is created for each photodiode element 14 and stored in the correction coefficient table storage unit 21 in the correction unit 20.

  When an actual CT image of the subject is taken, the output signals of the plurality of photodiode elements 14 are respectively measured by irradiating the subject with X-rays and detecting with a radiation detector (step 804). The measured output signal of the photodiode element 14 is transferred to the correction unit 20, and the correction unit 20 reads the correction coefficient corresponding to the received output signal from the table in the correction coefficient table storage unit 21, and outputs the correction signal to the output signal. Correction is performed by multiplication (step 805). The corrected output signal is output to the image processing unit of the CT apparatus and used to create a CT image (step 806).

  Factors that affect the energy characteristics of the photodiode element 14 include non-linear fluctuations caused by the incidence of X-rays on the edge of the scintillator element described in the present embodiment, the homogeneity of the scintillator element 11, and the detector module 101. There are mounting angles. As in steps 801 and 802 described above, by irradiating X-rays and measuring the output signal of the photodiode element 14 to obtain the correction coefficient, the correction coefficient can be obtained including the influence of factors other than those described above. . Therefore, a preferable CT image can be acquired.

  In addition, since the radiation detector of the present embodiment has a structure in which the collimator plate 16 and the shielding portion 17 are both bonded and fixed to the scintillator element array 13 by the adhesive layer 18, the collimator plate 16 is rotated along with the rotation of the X-ray CT apparatus. It is possible to prevent vibration. Therefore, it is possible to prevent an output error of the photodiode element 14 due to the position fluctuation of the collimator plate 16.

  Further, in the present embodiment, the collimator plate 16 can be formed into a simple plate shape using a heavy metal such as tungsten or molybdenum as in the prior art, so that it can be manufactured very simply and inexpensively. There are also benefits.

(Second embodiment)
FIG. 10 shows a detector module 101 according to the second embodiment of the present invention. The detector module of the second embodiment has the same structure as the detector module of FIG. 1, but differs from the first embodiment in that it does not include the collimator plate 16-r at the module end. Yes.

  Specifically, the reflecting material 12 is arranged on both end faces of the scintillator element array as in the first embodiment, but the collimator plate 16-r is arranged on one of the reflecting materials 12 on both end faces. Only the shielding part 17 is provided. When the radiation detector of the X-ray CT apparatus is configured by arranging the detector modules 101 in an arc shape as shown in FIG. 3 (b) by adopting a structure that does not include a collimator plate at one end in this way. Further, as shown in FIG. 11, the collimator plate 16-r does not physically interfere with the collimator plate 16-l of the adjacent detector module.

  As shown in FIG. 3 (b), when arranging detector modules as radiation detectors of an X-ray CT apparatus, a predetermined angle is set so that each detector module faces the focal point of the X-ray tube. Is arranged. All the collimator plates 16 are installed by changing the angle of the collimator plate 16 little by little so that the X-ray tube focal point is located on the extended line of the main plane. As a result, the collimator plate 16-r deleted in the second embodiment is directed to the X-ray tube focal point at substantially the same angle as the collimator plate 16-l of the adjacent detector module as shown in FIG. Will be. That is, only the collimator plate 16-l of the adjacent detector module can also serve as the collimator plate 16-r.

  As described above, even in a configuration that does not include the collimator plate 16-r on one end as in the second embodiment, the direct X-ray shielding effect is maintained as in the first embodiment. However, the function of the scattered radiation elimination collimator is not impaired. Moreover, according to the structure of the second embodiment, it is possible to avoid interference with adjacent detector modules.

(Third embodiment)
FIG. 12 shows a detector module according to the third embodiment of the present invention. This embodiment has the same structure as the detector module of FIG. 1 of the first embodiment, but the first embodiment is that the module end shields 17-l and 17-r are not provided. Is different. By not providing the shields 17-1 and 17-r at both ends, the width Wr-e of the reflector 12 at the end of the scintillator element array 13 is made smaller than the width Wr of the reflector 12 inside the scintillator element array 13. be able to. Thereby, when arrange | positioning a detector module in circular arc shape and comprising a radiation detector, it can prevent that adjacent modules interfere.

  In this embodiment, only the end portion is not provided with the shielding portion 17, but since the shielding portion 17 is provided between the scintillator elements 11 other than the end portion, the first detector module is the first. The substantially same effect as the embodiment can be obtained. Further, since the end shielding portion 17 is not provided, the assembling of the shielding portion can be facilitated.

(Fourth embodiment)
FIG. 13 shows a detector module 101 according to the fourth embodiment of the present invention. The present embodiment has the same structure as the detector module of FIG. 1 of the first embodiment, but does not include the collimator plate 16-r at one end of the detector module, and the shielding portions 17- l and 17-r are not provided. The width Wr-e of the reflecting material 12 at both ends is smaller than the width Wr of the reflecting material 12 inside the scintillator element array 13.

  The structure of this embodiment can obtain the same operations and effects as those of the second and third embodiments.

(Fifth embodiment)
Next, an X-ray CT apparatus provided with the radiation detector of the present invention will be described with reference to FIG.

  FIG. 14 is a block diagram showing an outline of the X-ray CT apparatus of the present invention. This apparatus includes a scan gantry unit 310 and an image reconstruction unit 320. The scan gantry unit 310 is attached to a rotating disk 311 having an opening 314 into which a subject is carried, an X-ray tube 312 that is a radiation source mounted on the rotating disk 311, and an X-ray tube 312. A collimator 313 that controls the radiation direction of the X-ray bundle, an X-ray detector 315 mounted on the rotating disk 311 facing the X-ray tube 312, and X-rays detected by the X-ray detector 315 as predetermined signals A detector circuit 316 for conversion and the like, and a scan control circuit 317 for controlling the rotation of the rotating disk 311 and the width of the X-ray bundle are provided. Any of the radiation detectors of the first to fourth embodiments is used for the X-ray detector 315.

  The image reconstruction unit 320 includes an input device 321 for inputting a subject name, examination date and time, examination conditions, and the like, and an image computation circuit 322 that performs CT image reconstruction by computing the measurement data S1 sent from the detector circuit 316. An image information adding unit 323 for adding information such as the subject name, examination date and time, and examination conditions input from the input device 321 to the CT image created by the image calculation circuit 322, and a CT image to which the image information is added And a display circuit 324 that adjusts the display gain of the signal S2 and outputs the adjusted signal S2 to the display monitor 330.

  In this X-ray CT apparatus, X-rays are irradiated from the X-ray tube 312 in a state where the subject is placed on a bed (not shown) installed in the opening 314 of the scan gantry unit 310. This X-ray obtains directivity by the collimator 313 and is detected by the X-ray detector 315. At this time, while rotating the rotating disk 311 around the subject, the direction of X-ray irradiation is changed. The X-ray transmitted through the subject is detected. A tomographic image created by the image reconstruction unit 320 based on the measurement data is displayed on the display monitor 330.

  The correction unit 20 can be arranged in the image calculation circuit 322 in addition to the arrangement arranged in the X-ray detection circuit 316.

  The radiation detector of the present invention has a high S / N ratio by preventing the generation of noise components caused by direct X-rays that have passed through the collimator plate entering the photodiode element array, and also prevents the deterioration of characteristics due to high-speed rotation. X-ray CT apparatus equipped with a scanner can be provided at low cost.

  The first to fifth embodiments described above are not intended to limit the structure of the present invention, but are examples showing specific embodiments, and other forms having the same effect. However, the present invention can be realized.

  11 Scintillator element, 12 Reflector, 13 Scintillator element array, 14 Photodiode element, 15 Photodiode element array, 16 Collimator plate, 17 Shielding part, 18 Adhesive layer, 101 Detector module, 102 Detector element module, 103 Scattered ray removal Collimator, 310 scan gantry, 311 rotating disk, 312 X-ray tube, 313 collimator, 314 aperture, 315 X-ray detector, 316 detector circuit, 317 scan control circuit, 320 image reconstruction unit, 321 input device, 322 Image calculation circuit, 323 image information adding unit, 324 display circuit, 330 display monitor, S1 measurement data, S2 CT image signal

The present invention relates to a radiation detector that detects X-rays, γ-rays, and the like, and more particularly to a medical image diagnostic apparatus such as an X-ray CT apparatus using the radiation detector.

In the structures disclosed in Patent Documents 1 and 2, the upper surface of the reflective material between the scintillator elements and the upper surface of the edge of the scintillator element are covered with a shielding member having a width wider than that of the reflective material. Not only can X-rays to be reduced be reduced, but also X-rays that pass through the scintillator element and enter the reflector from the side surface of the scintillator element can be reduced. However, since the edge of the scintillator element is covered with the shielding member, there is a problem that the aperture ratio of the scintillator element is substantially reduced and the amount of X-rays incident on the scintillator element is reduced. Further, in the techniques disclosed in Patent Documents 1 and 2, the collimator plate and the shielding member are separated, and when used in a high-speed X-ray CT apparatus of about 0.3 seconds / revolution in recent years, the collimator plate May vibrate. Patent Document 1 describes that the collimator plate may be connected to the shielding member, but there is no disclosure of a specific structure, and how the minute shielding member is connected to the collimator plate. Is not disclosed.

1 is a cross-sectional view of a detector module according to a first embodiment of the present invention. FIG. 2 is a block diagram showing a configuration in which a detector is provided in the detector module according to the first embodiment. (a) A perspective view of the detector module of FIG. 1, (b) a perspective view of the radiation detector of the first embodiment. The graph which shows the relationship between the dynamic range of the output of the photodiode element of the radiation detector for X-ray CT apparatuses of this invention, and the noise component current contained in an output signal current. (a) Explanatory diagram showing direct X-rays incident on the scintillator element 11 of the detector module of the first embodiment, (b) Explanatory diagram showing direct X-rays incident on the scintillator element 11 of the detector module of the comparative example . 3 is a graph showing an example of the output of the photodiode element 14 with respect to X-ray energy in the case of Ws <Wr and Ws> Wr, respectively. 6 is a graph showing fluctuations in the output of the photodiode element due to X-ray components incident on the edge portion of the scintillator element. 6 is a flowchart showing a procedure of energy characteristic correction according to the first embodiment. 6 is a graph showing the relationship between the object size and the output signal of the photodiode element in the detector module of the first embodiment. Sectional drawing of the detector module of 2nd Embodiment. FIG. 11 is a partial cross-sectional view when the detector modules of FIG. 10 are arranged side by side. Sectional drawing of the detector module of 3rd Embodiment. Sectional drawing of the detector module of 4th Embodiment. The block diagram of the X-ray CT apparatus of 5th Embodiment.

An adhesive layer 18 is disposed between the upper surface of the shielding part 17 and the end surface of the collimator plate 16. The upper surface of the shielding part 17 and the end surface of the collimator plate 16 are bonded by the adhesive layer 18. This prevents the collimator plate 16 from vibrating even when the radiation detector is used in an X-ray CT apparatus and rotated at a high speed, thereby preventing radiation detection errors due to fluctuations in the position of the collimator 16. Can do.

FIG. 4 is a graph showing the relationship between the dynamic range of the output of the photodiode element 14 of the radiation detector for the X-ray CT apparatus and the noise component current included in the output signal current, and the cause of the noise. As is clear from FIG. 4, in the region where the output signal current is extremely weak, the circuit noise is dominant, but in the region of the output signal current that is larger than a certain level, the quantum noise is dominant, and the CT image It can be seen that most of the noise is quantum noise. Therefore, in the present invention, the shielding unit 17 is designed to reduce quantum noise.

When an X-ray CT image of an actual subject is taken, the output signals of the plurality of photodiode elements 14 are measured by irradiating the subject with X-rays and detecting with a radiation detector (step 804). The measured output signal of the photodiode element 14 is transferred to the correction unit 20, and the correction unit 20 reads out the correction coefficient corresponding to the received output signal from the correction coefficient table in the correction coefficient table storage unit 21, and outputs it. Correction is performed by multiplying the signal (step 805). The corrected output signal is output to the image processing unit of the X-ray CT apparatus and used to create a CT image (step 806).

Claims (8)

A plurality of scintillator elements that are arranged to emit fluorescence by radiation, and a scintillator element array that is disposed at least between the adjacent scintillator elements and reflects the fluorescence; and
A plurality of photodiode elements that are arranged on a surface opposite to the radiation incident surface of the plurality of scintillator elements, and that detect the fluorescence emitted by the scintillator elements;
A plurality of collimator plates arranged in a space on the radiation incident side of the scintillator element array between the adjacent scintillator elements;
On the radiation incident surface of the scintillator element array, a shielding part that shields the reflecting material between the adjacent scintillator elements is arranged, and the width Ws of the shielding part is smaller than the width Wr of the reflecting material, Larger than the thickness Wc of the collimator plate,
An radiation layer is disposed between the upper surface of the shield and the end surface of the collimator plate, and the radiation detection is characterized in that the upper surface of the shield and the end surface of the collimator plate are adhered by the adhesive layer. vessel. The radiation detector according to claim 1, further comprising a correction unit that corrects the output of the photodiode element,
The correction unit generates an output signal of the photodiode element for each of the plurality of scintillator elements resulting from the fact that the radiation that has passed between the collimator plates is incident on the edge of the scintillator element and converted into fluorescence. A radiation detector characterized by correcting variations that occur.   3. The radiation detector according to claim 2, wherein the correction unit uses a correction coefficient predetermined for each scintillator element to correct the variation, and outputs an output signal of the photodiode element for each scintillator element. A radiation detector characterized by correcting the above.   4. The radiation detector according to claim 3, wherein the correction coefficient is prepared for each energy of the radiation.   2. The radiation detector according to claim 1, wherein the adhesive layer covers the shielding part, and the shielding part is adhered to the scintillator element array.   2. The radiation detector according to claim 1, wherein the reflector is disposed on both end faces of the scintillator element array, and the collimator plate is not disposed on one of the reflectors on the both end faces. Radiation detector.   2. The radiation detector according to claim 1, wherein the reflector is disposed on both end faces of the scintillator element array, and the width of the reflector on the both end faces is the width of the reflector located between the scintillator elements. The radiation detector is characterized in that it is smaller than Wr, and the shielding portion is not disposed on at least one of the reflecting materials on both end faces. A radiation source, a radiation detector disposed opposite to the radiation source, a rotating disk that holds the radiation source and the radiation detector and is driven to rotate around a subject, and is detected by the radiation detector In an X-ray CT apparatus comprising an image reconstruction unit that reconstructs a tomographic image of the subject based on the intensity of the received radiation,
An X-ray CT apparatus using the radiation detector according to claim 1 as the radiation detector.

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