JP2004184163A - Radiation detector and medical image diagnostic apparatus - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector element array having a small fluctuation of a light emission property in a block so as to obtain a CT image with less artifact. <P>SOLUTION: In the radiation detector having a plurality of fluorescent elements one-dimensionally or two-dimensionally arrayed on a photodiode array substrate, an orientation defined by the fluorescent element (a scintillator element) and reflection layers disposed at its both sides is the same as an incident direction of the radiation and an angle formed by a contact face of the fluorescent element and the reflection layer and a substrate face changes in the direction from the center of the substrate to an end. The fluctuation of the light emission property of the radiation radially entering in the detector block is suppressed. The sensitive radiation detector can be provided. Since such radiation detector is provided, a medical image diagnostic apparatus such as an X-ray CT apparatus with less artifact can be provided. <P>COPYRIGHT: (C)2004,JPO&NCIPI

Description

[0001]
TECHNICAL FIELD OF THE INVENTION
The present invention relates to a radiation detector for detecting X-rays and γ-rays, and more particularly to a radiation detector suitable for a medical image diagnostic device such as an X-ray CT device. Further, the present invention relates to a medical image diagnostic apparatus using the radiation detector.
[0002]
[Prior art]
2. Description of the Related Art As a radiation detector used in a medical image diagnostic apparatus such as an X-ray CT apparatus, a solid state detector in which a phosphor element and a photodiode are combined is widely used. The structure of a solid state detector is generally a combination of a scintillator that absorbs X-rays and emits visible light, and a photodiode that converts the emitted light into an electric signal. These detector elements are one-dimensional. Alternatively, two-dimensionally arranged element arrays are used. In such a detector element array, it is necessary to separate the scintillator into channels in order to prevent crosstalk between adjacent channels, and a reflection layer is usually formed between adjacent scintillator elements. In the case of a one-dimensional array detector, for example, the scintillator elements and the reflective layer have a rectangular parallelepiped shape in which a channel-separated scintillator element has a width of 0.9 mm, a length of 30 mm, and a height of 1.8 mm. The reflection layer provided between the elements has, for example, a rectangular parallelepiped shape of 0.1 mm in width × 30 mm in length × 1.8 mm in height. Such a scintillator element and a reflective layer alternately arranged are adhered on a plane of a photodiode array substrate to constitute a detector element array block. The structure of each detector element is described in Patent Document 1, for example.
[Patent Document 1] JP-A-5-256949
[0003]
In the X-ray CT apparatus, a plurality of such detector element array blocks are arranged on an arc centered on the focal point of the X-ray tube to form a detector element array of, for example, 960 channels in total. At this time, in order to prevent scattered X-rays scattered by the subject from being incident on the detector element, the solid state detector is provided with a scattered radiation removing collimator. The plurality of scattered radiation removing collimators are aligned with the detector element array with high precision so as to cover the gap between the detector elements, and are arranged in an arc shape.
[0004]
[Problems to be solved by the invention]
As described above, in the conventional X-ray CT apparatus, the scattered radiation removing collimator is arranged on an arc, whereas the detector element array is on a polygonal pseudo-arc because the photodiode array substrate is flat. Are arranged. FIG. 8 shows a conventional detector element array, and FIG. 9 shows the relationship between such a detector element array and a scattered radiation removing collimator. In FIG. 8, the solid arrows indicate the direction of the focal point, and the dotted arrows indicate the normal direction of the substrate.
[0005]
As is clear from the figure, when focusing on one detector element array block 90, the scintillator element 91 at the center of the block is oriented toward the X-ray tube focal point, but the scintillator element 92 at the end of the block is That is, it does not face the direction of the X-ray tube focal point. As a result, the scintillator element 92 and the scattered radiation removing collimator 93 are shifted from each other toward the end of the element array block, and the apparent irradiation area from the focal point of the X-ray tube becomes wider. For this reason, the amount of light emission at the time of X-ray irradiation increases, and the output changes. Further, since the path of the X-ray passing through the element differs between the scintillator element at the center of the block and the scintillator element at the end, the X-ray energy characteristics also change. For these reasons, variations occur in the light emission characteristics in the element array block, and artifacts tend to occur in the CT image.
[0006]
Generally, the sensitivity variation in the element array block can be easily corrected by software processing during image processing, but the variation in the X-ray energy characteristics causes the output ratio to vary when the tube voltage of the X-ray tube is changed. This is a phenomenon that is difficult to correct by software processing.
[0007]
Such a situation is the same in a radiation detector having a two-dimensional detector element array. For example, as the element array arranged in the direction orthogonal to the rotation direction of the rotating plate on which the X-ray tube and the radiation detector are mounted, that is, in the direction of the rotation axis, the X-rays move away from the position (center) facing the X-ray tube. They are arranged in a direction deviating from the direction of the focal point of the tube, so that artifacts due to variations in X-ray energy dependence in the body axis direction are likely to occur.
[0008]
As a solution to these problems, it is conceivable to increase the thickness of the scattered radiation removing collimator. In that case, however, the dose of the X-rays incident on the solid state detector is reduced, resulting in a reduction in S / N. There is.
Therefore, an object of the present invention is to provide a radiation detector that suppresses variations in output in the arrangement direction of elements. Another object of the present invention is to provide a medical image diagnostic apparatus such as an X-ray CT apparatus or the like that can provide high image quality by reducing artifacts due to output variations by providing such a radiation detector.
[0009]
[Means for Solving the Problems]
In order to solve the above-described problems, the radiation detector of the present invention includes a plurality of phosphor elements arranged on a substrate, wherein the orientation of the phosphor elements determined by the phosphor elements and the reflective layers disposed on both sides of the phosphor elements is radiation. It is arranged so as to face the source (focal point).
[0010]
That is, the radiation detector of the present invention comprises a substrate, a plurality of phosphor elements arranged on the substrate, and a reflection layer disposed between adjacent phosphor elements to prevent crosstalk between adjacent phosphor elements. And a radiation detector including a plurality of photoelectric conversion elements provided at positions on the substrate corresponding to the phosphor elements,
The angle formed between the contact surface between the phosphor element and the reflective layer and the substrate surface at the center of the substrate changes from the center to the end of the substrate along the arrangement direction of the phosphor elements. Things.
[0011]
According to this radiation detector, for all the phosphor elements arranged in a plane, the orientation of the phosphor element defined by the phosphor element and the reflective layers on both sides thereof is substantially the same as the incident direction of radiation. Therefore, the apparent irradiation area of the radially incident radiation and the passage in the element can be made uniform, and the variation in the output of the element in the arrangement direction can be eliminated.
[0012]
The angle formed between the contact surface between the phosphor element and the reflective layer and the substrate surface at the center of the substrate (the angle on the substrate center side) is, for example, the phosphor element located at the center of the substrate or at the position closest to the center. The angle between the contact surface of the phosphor and the reflective layer is 90 °, and from the angle of 90 ° to -θ (θ = tan^-1(Ld / Lp), Ld is the distance from the radiation source to the upper surface of the phosphor element, and Lp is the pitch of the phosphor elements. In consideration of the processing accuracy of the radiation detector and the like, an error of ± 0.2 ° can be allowed.
[0013]
That is, when the number of phosphor elements in the arrangement direction is an even number, the angle between the contact surface between the reflective layer and the phosphor element located on the edge side of the substrate and the substrate surface at the central portion of the substrate (the central portion of the substrate). Angle) in order from the center of the substrate_Rn(N = 0, 1, 2,...), Θ_Rn= 90−a (n · θ) ± 0.2 degrees (0 <a ≦ 1, n = 0, 1, 2,...). When the number of phosphor elements is an odd number, the angle between the contact surface between the base phosphor element and the reflective layer located at the edge of the substrate and the substrate surface at the center of the substrate (the angle at the center of the substrate) is determined. , In order from the substrate center_Sn(N = 0, 1, 2,...), Θ_Sn= 90−a (n · θ) ± 0.2 degrees (0 <a ≦ 1, n = 0, 1, 2,...).
[0014]
In the radiation detector of the present invention, by changing the direction of the phosphor elements in the arrangement direction, the width changes along the depth direction. This change in width can be imparted to both or one of the phosphor element and the reflective layer. In the present specification, the width of the phosphor element and the reflection layer means the width in the arrangement direction, and the distance from the radiation incidence surface to the substrate surface is called height or depth.
[0015]
A medical image diagnostic apparatus according to the present invention includes a radiation source, a radiation detector disposed opposite to the radiation source with a subject interposed therebetween, and includes the above-described radiation detector as a radiation detector. . A particularly preferred medical image diagnostic apparatus of the present invention is a radiation source, a radiation detector arranged opposite to the radiation source, and holds the radiation source and the radiation detector, and is driven to rotate around the subject. Rotating disk, and image reconstruction means for reconstructing an image of the tomographic image of the subject based on the intensity of radiation detected by the radiation detector, using the radiation detector of the present invention as the radiation detector Things.
[0016]
In this medical image diagnostic apparatus, a radiation detector having a reduced output variation in the array direction is employed, so that artifacts due to the output variation, particularly the variation in radiation energy dependence, which is difficult to correct by software, are considered. Can be effectively suppressed.
[0017]
BEST MODE FOR CARRYING OUT THE INVENTION
Hereinafter, embodiments of the present invention will be described with reference to the drawings.
[0018]
First, as one embodiment of the radiation detector of the present invention, a radiation detector arranged in a one-dimensional direction by combining a plurality of scintillators (phosphor elements) and photodiodes (photoelectric conversion elements) will be described. As shown in FIG. 1, the radiation detector 10 includes a scintillator block in which a plurality of scintillator elements 12 and a reflective layer 13 are alternately stacked on a flat substrate 11 so that the stacking direction is parallel to the substrate surface. A photodiode 14 is formed in a portion corresponding to each scintillator element of the substrate 11 in the same arrangement as the arrangement of the scintillator elements 12. As a result, a scintillator element surrounded by the reflective layer 23 and channel-separated, and a photodiode below the scintillator element constitute one detection element.
[0019]
The size of the scintillator element and the reflective layer and the number of scintillator elements arranged depend on the use of the radiation detector and are not particularly limited. For example, in the case of a detector array for X-ray CT, for example, about 20 to 40 mm wide substrate The scintillator elements 12 are provided at an arrangement pitch of 1 mm. The width of the reflective layer 13 in the arrangement direction is usually about 1/10 to 2/10 of the arrangement pitch.
[0020]
The radiation detector of the present invention is configured such that the angle formed between the contact surface between the scintillator element 12 and the reflective layer 13 and the substrate surface gradually changes while keeping the arrangement pitch of the scintillator elements constant. The change in the angle is such that the direction of the normal to the contact surface between the scintillator element 12 and the reflective layer 13 is substantially perpendicular to the line connecting the scintillator element 12 and the radiation source (if X-ray, its focal point). To
[0021]
If the number of scintillator elements provided on the substrate is even, the center of the reflective layer 13 is located at the center of the substrate 11, as shown in FIG. In this case, for example, the angle θ between the contact surfaces of the scintillator elements 121 on both sides in contact with the reflection layer 13 and the substrate plane_R0Is 90 °, and the reflection layer on both sides of these scintillator elements 121, 121 has an angle θ between the contact surface with the scintillator element on the end side thereof and the substrate surface._R1(Angle at the center of the substrate) is 90 ° or less. Thereafter, the angle between the contact surface and the substrate surface is made smaller as approaching the end.
[0022]
When the number of scintillator elements provided on the substrate is odd, the center of the scintillator element 120 is located at the center of the substrate 11 as shown in FIG. In this case, for example, the angle θ between the contact surfaces of the reflective layers 131 and 131 on both sides in contact with the scintillator element 120 and the substrate plane_S0Is 90 °, and the angle θ between the contact surface between the scintillator elements 121 and 121 on the end side and the reflection layers 131 and 131 on the end side with respect to the substrate surface is set to 90 °._S1(Angle at the center of the substrate) is 90 ° or less. Thereafter, the angle between the contact surface and the substrate surface is made smaller as approaching the end.
[0023]
As described above, the angle is changed so that the direction of the scintillator element matches the direction of the radiation source. Generally speaking, when the number of phosphor elements in the arrangement direction is an even number, the contact surface between the reflective layer 13 and the scintillator element 12 located on the edge side of the substrate and the substrate surface at the central portion of the substrate are Angle (the angle at the center of the substrate) in the order from the center of the substrate to θ_Rn(N = 0, 1, 2, ...)_,
θ_Rn= 90-a (n · θ) ± 0.2 degrees
(Where 0 <a ≦ 1, n = 0, 1, 2,..., Θ = tan^-1(Lp / Ld), Ld is the distance from the radiation source to the upper surface of the scintillator element, and Lp is the arrangement pitch of the scintillator element. The same applies hereinafter). When the number of scintillator elements is an odd number, the angle between the contact surface between the scintillator element 12 and the reflective layer 13 located on the side of the substrate and the substrate surface at the substrate center (the angle on the substrate central side) From the center of the substrate_Sn(N = 0, 1, 2,...)
θ_Sn= 90-a (n · θ) ± 0.2 degrees
Becomes
[0024]
θ (= tan^-1(Lp / Ld)) corresponds to the elevation angle per pitch of the scintillator element. Thus, by inclining the direction of the scintillator element by the elevation angle for each pitch, the scintillator element always faces the radiation source. That is, it is possible to prevent the radiation incident on the adjacent scintillator element from entering, and it is possible to suppress the variation in the amount of emitted light.
[0025]
In the equation, “a” defines an allowable inclination angle for obtaining the effect of the present invention, and ± 0.2 degrees is an allowable error in consideration of a dimensional error in manufacturing. That is, according to the study of the present inventors, the effect of the present invention can be obtained if the element array block is manufactured with an inclination angle of each element of ± 0.2 degrees, but the variation is larger. If so, the effect of the present invention is reduced. Also, the inclination angle of each element is 0 <θ._nThe effect of the present invention can be obtained in the range of ≦ n · θ, but when the inclination is more than that, the light emission characteristics deteriorate. Therefore, the angle between the contact surface (boundary surface) of each element and the substrate surface is θ_n= 90−a (n · θ) ± 0.2 degrees (0 <a ≦ 1, n = 0, 1, 2,...).
[0026]
The effect of the present invention will be described more specifically. For example, as shown in FIG. 3, the arrangement pitch of the scintillator elements (the width of the pair of the scintillator element and the reflective layer) Lp is 1 mm on an arc having a radius of 1037 mm at a distance of Ld = 1037 mm from the focal point of the X-ray tube 31. It is assumed that a 24-channel detector element array block 32 is arranged. Here, in the case where the detector element array block 32 is a conventional detector element array block 80 including a scintillator element 81 and a reflective layer 82 having a rectangular cross-sectional shape as shown in FIG. Is arranged in parallel with the normal direction of the substrate, and thus faces in the direction of the focal point of the X-ray tube. However, as the reflection layer other than the center of the substrate moves away from the center of the substrate, θ (= tan^-1(Lp / Ld)) = approximately 0.055 degrees from the direction of the X-ray tube focal point, and the reflection layer at the end of the block is displaced by approximately 0.66 degrees (= 12 × θ) from the direction of the X-ray focal point. Will be. That is, the angle between the normal to the side of the reflective layer at the center of the substrate (the surface intersecting with the substrate surface) and the direction of the focal point of the X-ray tube is 90 degrees, and the normal to the side of the nth reflective layer from the center of the substrate. And the direction of the focal point of the X-ray tube (the complementary angle in the case of an obtuse angle) is θ_Rn= 90−n · θ, and at the end of the block, 90−0.66 = about 89.34 degrees. Naturally, the scintillator element is similarly shifted from the direction of the X-ray tube focal point. The same applies to the case where an odd number of scintillator elements are arranged. The scintillator element in the center of the block faces the direction of the X-ray tube focal point. It deviates from the direction of the tube focus. Therefore, the angle between the normal to the side of the scintillator element and the direction of the X-ray tube focal point is θ_Sn= 90−n · θ (n = 0, 1, 2,...).
[0027]
On the other hand, in the case of the detector element array block shown in FIG. 1, in the present invention, the angle between the side surface of the reflective layer and the substrate surface (the angle on the substrate central portion side) θ_Rn(N = 0, 1, 2, ...) is θ_Rn= 90−n · θ (n = 0, 1, 2,...), The scintillator elements 12 and the reflection layers 13 are arranged. Therefore, even when the element array blocks are arranged on a polygonal pseudo-arc. , All the elements can be arranged in the direction of the X-ray tube focal point.
[0028]
Although FIGS. 1 and 2 show the case where the width of the reflective layer is constant in the depth direction, the width of the scintillator element may be constant in the depth direction as shown in FIG. Although only one-dimensional arrangement is shown in the drawing, a two-dimensional element array may have a similar structure in each arrangement direction.
[0029]
Next, a method for manufacturing a radiation detector having such a configuration will be described.
Generally, a method of manufacturing a channel-separated scintillator element block is as follows. 1) After laminating a plurality of plate-shaped scintillator materials at a predetermined interval via a reflective material to form a block, the laminated block is cut into a predetermined thickness. And 2) bonding the plate-shaped scintillator material on the photodiode array substrate, and forming cut grooves at a predetermined pitch in the plate-shaped scintillator material. There is a method of adding a material constituting the reflective layer. Further, as a method proposed by the present applicant (Japanese Patent Application No. 2002-162587), 3) cut grooves are formed at a predetermined pitch in a plate-shaped scintillator material, and reflection is performed in the grooves. After filling the material constituting the layer by a specific method, it is cut into a predetermined thickness, and this is cut on the substrate on which the photodiode array is formed. There is a method of bonding. The radiation detector of the present invention can employ any of the above manufacturing methods, but the method 3) is particularly preferable in terms of ease of manufacture and performance of the obtained element.
[0030]
FIG. 5 shows an example of a method for manufacturing scintillators arranged two-dimensionally. In this method, first, a cut groove 51 having a predetermined pitch, a predetermined width, and a predetermined depth is formed in a rectangular parallelepiped scintillator material 50 while leaving the bottom portion not to be cut off. As the scintillator material, a phosphor having a high radiation-to-light conversion efficiency such as a rare earth phosphor, for example, Ce as a light emitting element described in PCT / JP98 / 05806 or Japanese Patent Application No. 11-176719, and at least Gd, Al , A rare earth oxide scintillator made of a garnet-based host crystal containing Ga, O, or the like can be used. Next, the cut groove 51 is filled with a material 52 for forming a reflective layer. If the material 52 is a material that does not cure and shrink, the cut groove 51 has a predetermined angle with respect to the scintillator material surface. It is formed as follows. This angle is such that, when a detector element array is formed, the angle formed by the contact surface between the reflective layer and the scintillator element and the substrate surface is inclined from the center to the end of the substrate. This can be achieved by sequentially changing the angle between the material and the cutting blade in a processing machine such as a wire saw or a slicer from the center to the end. On the other hand, if the material 52 is a material that cures and contracts, the cut grooves 51 may be formed in parallel.
[0031]
After the cut grooves 51 are formed in this manner, a material 52 for forming a reflective layer is filled in the grooves. As the material 52 for forming the reflection layer, a material in which an inorganic compound powder such as titanium oxide, aluminum oxide, barium sulfate or the like is contained as a reflection material in an organic resin such as an epoxy resin, an acrylic resin, or a phenol resin, or a light reflection material A high-efficiency metal plate, a combination thereof, or the like can be used. Here, in the case where a material containing a reflecting material in a resin is used as the material 52, centrifugal injection is performed using a centrifuge. As a result, the material 52 can be injected to the bottom of the groove, and air bubbles in the resin can be substantially eliminated.
[0032]
Next, when a material in which a reflective material is dispersed in an organic resin is used as the material 52, the organic resin in the material 52 is cured to form a reflective layer. When a metal plate or the like is inserted as it is or through an adhesive, the adhesive is cured as necessary. Thereafter, the slice is sliced to a desired thickness in a direction orthogonal to the depth direction of the groove. Here, when a material in which a reflecting material is dispersed in an organic resin is used as the material 52, the material shrinks during curing, but the bottom of the groove does not shrink because it is fixed to the bottom of the scintillator material. A slope is formed from the bottom to the top. The degree of cure shrinkage varies depending on the type and viscosity of the resin, and by appropriately selecting these conditions, a desired inclination can be imparted to the groove. For example, when an epoxy resin having a viscosity of 10,000 to 20,000 cPs is used as the organic resin, an inclination of about 0.04 to 0.06 degrees can be formed for a groove having a width of 0.15 mm and a depth of 10 mm. Therefore, in the notch groove forming step, by forming grooves parallel to the depth direction and using a material 52 containing an organic resin appropriately selected, the above-described desired angle can be formed in the reflective layer.
[0033]
On the other hand, when a material without such curing shrinkage, for example, a metal plate or the like is used, a reflection layer having a desired angle can be formed by inclining the cut groove 51 itself in advance.
The radiation detector of the present invention is obtained by adhering the scintillator thus separated into channels in the same arrangement pitch as the arrangement pitch of the scintillator elements on the substrate 53 on which a photodiode array is formed in advance. As the photodiode, a known photodiode such as a PIN photodiode can be used.
[0034]
Next, an X-ray CT apparatus using the radiation detector of the present invention will be described. FIG. 6 is a diagram schematically showing an X-ray CT apparatus as an example of medical image diagnosis to which the present invention is applied.
[0035]
The X-ray CT apparatus includes a scan gantry unit 610 and an image reconstruction unit 620. The scan gantry unit 610 includes a rotating disk 611 having an opening 614 into which a subject is loaded, and the rotating disk 611. An X-ray tube 612 mounted on the X-ray tube 611, a collimator 613 mounted on the X-ray tube 612 and controlling the radiation direction of the X-ray flux, and an X-ray detector mounted on the rotating disk 611 facing the X-ray tube 612 A detector 615, a detector circuit 616 for converting the X-rays detected by the X-ray detector 615 into a predetermined signal, and a scan control circuit 617 for controlling the rotation of the rotating disk 611 and the width of the X-ray flux. ing.
[0036]
The image reconstructing unit 620 is an input device 621 for inputting the subject's name, examination date and time, examination conditions, and the like, and an image computing circuit that performs computational processing on the measurement data S1 sent from the detector circuit 616 to reconstruct a CT image. 622, an image information adding unit 623 for adding information such as a subject name, an examination date and time, and examination conditions input from the input device 621 to the CT image created by the image arithmetic circuit 622; A display circuit 624 that adjusts the display gain of the CT image signal S2 and outputs the adjusted display gain to the display monitor 630.
[0037]
In this X-ray CT apparatus, X-rays are emitted from an X-ray tube 612 in a state where a subject is placed on a bed (not shown) installed in an opening 614 of a scan gantry section 610. The X-ray has directivity by a collimator 613 and is detected by an X-ray detector 615. At this time, by rotating the rotating disk 611 around the subject, X-rays transmitted through the subject are detected while changing the direction of X-ray irradiation. The tomographic image created by the image reconstruction unit 620 based on the measurement data is displayed on the display monitor 630.
[0038]
The X-ray detector 615 is configured by arranging a large number of detector element array blocks in which a plurality of detection elements each of which is a combination of a scintillator element (phosphor element) and a photodiode (photoelectric conversion element) are arranged in an arc shape. The radiation detector of the present invention as shown in FIG. 1, FIG. 2 or FIG. 4 is used as this detector element array block. The radiation detector may be a one-dimensional detector in which the elements are arranged in a row in the circumferential direction, or a two-dimensional detector in which several elements are arranged in a row. In the case of a two-dimensional detector, each element is also arranged in the direction of the X-ray tube focal point in the arrangement direction parallel to the rotation axis of the rotating disk 611 (the body axis direction of the subject).
[0039]
The X-ray detector 615 is arranged on an arc centered on the focal point of the X-ray tube 612 so that the side where the scintillator element is arranged faces the radiation source. Here, the X-ray detector 615 is located at the end of the block because the direction of the scintillator element determined by the scintillator element and the reflective layers on both sides of the scintillator element coincides with the X-ray irradiation direction from the focal point. Even in the case of a scintillator element, radiation is incident similarly to the scintillator element at the center, and a uniform output can be obtained.
[0040]
FIG. 7 schematically shows X-ray energy characteristics when the conventional X-ray detector is used (b) and when the X-ray detector of the present invention is used (a). As shown, in the conventional X-ray detector, the X-ray energy characteristics in the element array block vary greatly. As a result, variations in light emission characteristics occur, and artifacts tend to occur in CT images. On the other hand, in the X-ray detector of the present invention, variation in X-ray energy characteristics inside the block is suppressed, and variation in light emission characteristics can be reduced, so that a high-quality CT image with less artifact can be obtained. it can.
[0041]
【Example】
Hereinafter, embodiments of the radiation detector of the present invention will be described.
[0042]
<Example 1>
Gd₃(Al, Ga)₅O₁₂: XY plane of a scintillator plate (having dimensions of 26 × 30 × 1.8 mm in X direction, Y direction and Z direction each) composed of a Ce ceramic scintillator is provided with a 1-mm pitch, 24 channel one-dimensional photodiode array. It adhered to the formed substrate. Then, using a diamond grindstone having a thickness of 0.13 mm, a 1.8 mm deep groove was formed in the scintillator plate at a pitch of 1 mm in the X direction parallel to the Y direction. At this time, when processing the groove closest to one end of the scintillator plate, it was set so that the grindstone entered at an angle of 0.66 degrees with respect to the Z direction (the normal direction of the photodiode array substrate). The second and subsequent grooves from the end are grooved so that the angle of the grindstone gradually decreases by 0.055 degrees so that the angle of the grindstone becomes 0 degree at the center groove (13th groove) of the scintillator plate. Was done. The groove on the other end side from the center was processed such that the angle of the grindstone was gradually increased in the opposite direction, and the groove (25th) closest to the other end was 0.66 degrees. Thus, a 24-channel element array was formed.
[0043]
After the groove is formed by changing the inclination angle in this manner, TiO is added to the groove.₂An epoxy resin mixed with the powder was poured and cured to form a reflective layer, thereby producing a detector element block. In this detector element block, after bonding the scintillator material to the substrate, the groove serving as the reflective layer was processed with a grindstone having a constant thickness. Therefore, as shown in FIG. 2, both sides of the reflective layer had a parallel shape. Was.
[0044]
Forty detector element blocks thus manufactured were mounted on an X-ray CT apparatus, and the emission characteristics (X-ray energy characteristics) and variations thereof were measured. The distance from the focal point of the X-ray tube to the detector element block was 1037 mm. The X-ray energy characteristic E was defined by the following formula according to the output ratio of each channel at the time of X-ray irradiation at a tube voltage of 100 kV and 120 kV and a tube current of 100 mA.
E = 6095 · log ｛(R₁₀₀/ X₁₀₀) / (R₁₂₀/ X₁₂₀)｝
Where X₁₀₀(X₁₂₀), R₁₀₀(R₁₂₀) Are sample and reference outputs at a tube voltage of 100 kV (120 kV), respectively.
[0045]
The variation in the X-ray energy characteristics was defined by the difference between the maximum value and the minimum value of the characteristic E of each channel in the detector element block. If this variation is 5 or more, a ring artifact tends to occur in the CT image.
The variation of the X-ray energy characteristics within the block of this embodiment was about 0.3, where the variation of the conventional element block was 1. Therefore, it was shown that this X-ray CT apparatus can suppress CT image artifacts.
[0046]
<Example 2>
Gd₃(Al, Ga)₅O₁₂: Using a multi-wire saw having a wire diameter of 0.13 mm on the XY plane of a scintillator plate (each dimension in the X, Y and Z directions is 24 × 36 × 4 mm) composed of a Ce ceramic scintillator And a grid-like grooving process parallel to the Y direction was performed. In the Y direction, 23 grooves were machined at a pitch of 1 mm in the X direction, and in the X direction, 15 grooves were machined at a pitch of 2.25 mm in the Y direction. Each groove had a depth of 3 mm, and was connected at the bottom of the scintillator plate. Each groove was machined vertically to the bottom of the scintillator.
[0047]
The groove of the scintillator plate having been subjected to the groove processing in this manner is provided with TiO.₂An epoxy resin mixed with the powder was poured and cured to form a reflective layer. At this time, curing shrinkage of the epoxy resin occurs. However, since each scintillator element is connected at the bottom, it has a shape inclined toward the center. Thereafter, the upper and lower surfaces were cut and polished to a thickness of 1.8 mm, thereby producing a 24 × 16 channel scintillator element array. In this scintillator element array, the scintillator element was processed with a wire saw, so that both side surfaces were parallel, but the reflective layer had a shape in which the upper surface side in the Z direction became thin due to curing shrinkage. The angle formed between the side surface of the scintillator element and the substrate surface was 89.43 degrees for the end element in the X direction and 89.14 degrees for the end element in the Y direction.
[0048]
The scintillator element array produced as described above was adhered to a 24 × 16 channel photodiode array substrate to obtain a detector element block. This detector element block was mounted on an X-ray CT apparatus so that the X direction was the circumferential direction, and the light emission characteristics were measured as in Example 1. As a result, when the variation in the conventional element block in both the X direction and the Y direction is 1, the value is about 0.4, which indicates that the X-ray CT apparatus employing this detector element block can suppress CT image artifacts. Was done.
[0049]
【The invention's effect】
According to the present invention, in a radiation detector in which a plurality of phosphor elements are arranged in a one-dimensional or two-dimensional direction, a phosphor element defined by a phosphor element (scintillator element) and reflection layers disposed on both sides thereof The direction of the radiation is substantially the same as the incident direction of the radiation, thereby suppressing variation in the light emission characteristics in the detector block with respect to the radiation incident radially, and providing a highly accurate radiation detector. be able to. Further, by providing such a radiation detector, a medical image diagnostic apparatus such as an X-ray CT apparatus with less artifact can be provided.
[Brief description of the drawings]
FIG. 1 is a diagram showing a cross section of an embodiment of the radiation detector of the present invention.
FIG. 2 is a diagram showing a cross section of a main part of the radiation detector element array block of FIG. 1;
FIG. 3 is a diagram showing a relationship between a radiation source and a radiation detector element array block.
FIG. 4 is a diagram showing a cross section of another embodiment of the radiation detector of the present invention.
FIG. 5 is a diagram showing an example of a method for manufacturing a radiation detector of the present invention.
FIG. 6 is a diagram showing a configuration of an X-ray CT apparatus equipped with the radiation detector of the present invention.
FIG. 7 is a diagram showing X-ray energy characteristics of the radiation detector of the present invention and a conventional radiation detector.
FIG. 8 is a diagram showing a cross section of a conventional radiation detector element array block.
FIG. 9 is a diagram showing a relationship between a conventional radiation detector element array and a scattered radiation removing collimator.
[Explanation of symbols]
11: scintillator element, 12: reflective layer, 13: photodiode, 14: photodiode array substrate, 31: X-ray tube, 32: detector element array block, 610 ..Scan gantry part, 611: rotating disk, 612: X-ray tube, 613: collimator, 614: opening, 615: X-ray detector, 616: detector Circuit, 617: scan control circuit, 620: image reconstruction unit, 621: input device, 622: image operation circuit, 623: image information addition unit, 624: display circuit, 630 ・ ・ ・ Display monitor

Claims (5)

A substrate, a plurality of phosphor elements arranged on the substrate, a reflection layer disposed between adjacent phosphor elements to prevent crosstalk between adjacent phosphor elements, and a phosphor element of the substrate. In a radiation detector including a plurality of photoelectric conversion elements provided at corresponding positions,
A radiation detector, wherein the angle formed between the contact surface between the phosphor element and the reflective layer and the substrate surface changes from the center to the end of the substrate along the arrangement direction of the phosphor elements. 2. The radiation detector according to claim 1, wherein a distance from the radiation source to the upper surface of the phosphor element is Ld, an arrangement pitch of the phosphor elements is Lp, and θ = tan ⁻¹ (Lp / Ld). When the number of arrays is even, the angle between the contact surface between the reflective layer and the phosphor element located at the edge of the substrate and the substrate surface (the angle at the substrate central portion) is θ _{R n in} order from the substrate center. (N = 0, 1, 2,...), Θ _{R n} = 90−a (n · θ) ± 0.2 degrees (0 <a ≦ 1, n = 0, 1, 2,...) When the number of phosphor elements arranged is an odd number, the angle between the contact surface between the phosphor element and the reflective layer located at the end of the substrate and the substrate surface (the angle at the center of the substrate) is determined in order from the center of the substrate θ _Sn (n = 0,1,2 ···) to the _{θ Sn = 90-a (n} · θ) ± 0.2 degrees (0 <a ≦ 1 A radiation detector, characterized in that n = 0, 1, 2 ...). 3. The radiation detector according to claim 1, wherein each of the phosphor elements has a constant width in an arrangement direction in a depth direction. 4. 3. The radiation detector according to claim 1, wherein each of the reflection layers has a constant width in an arrangement direction in a depth direction. 4. A radiation source, a radiation detector arranged opposite to the radiation source, a rotating disk holding the radiation source and the radiation detector and driven to rotate around a subject, and detecting the radiation with the radiation detector 5. The medical image diagnostic apparatus according to claim 1, further comprising an image reconstructing unit configured to reconstruct an image of the tomographic image of the subject based on the intensity of the obtained radiation. 6. The radiation according to claim 1, wherein the radiation detector is used as the radiation detector. A medical image diagnostic apparatus using a detector.

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