WO2004051312A1 - Radiation detector and medical image diagnosing device - Google Patents

Abstract

A radiation detector comprising a plurality of phosphor elements arranged on a photodiode array substrate in a one-dimensional or two-dimensional direction, wherein an angle formed by the contact surface between a phosphor element and a reflection layer and a substrate surface is allowed to vary from the center toward the end of the substrate so that the direction of a phosphor element defined by a phosphor element (scintillator element)and reflection layers on the opposite sides thereof is substantially identical to the radiation incident direction. Accordingly, variations in light emission characteristics, in a detection block, of a radiation being beamed radially are restricted to provide a high-accuracy radiation detector. The provision of such a radiation detector can offer a medical image diagnosing device such as an X-ray CT device with a minimum artifact.

Description

 Radiation detectors and medical diagnostic imaging equipment 'Technical field'

 The present invention relates to a radiation detector that detects X-rays, γ-rays, and the like, and particularly relates to a radiation detector suitable for a medical image diagnostic device such as an X-ray CT device or PET. The present invention also relates to a medical image diagnostic device using the radiation detector. Background technology

 Solid-state detectors, which combine a phosphor element and a photodiode, are widely used as radiation detectors for medical image diagnostic devices such as X-ray CT and PET. The structure of a solid state detector is generally a combination of a scintillator that emits visible light by absorbing X-rays and a photodiode that converts the emitted light into an electric signal. Element arrays arranged in one or two dimensions are used. In such a detection element array, the scintillator needs to be channel-separated 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 reflection layer have a rectangular parallelepiped shape for a channel-separated scintillator element, and the reflection layer provided between the elements has a rectangular parallelepiped shape. Such a scintillator element and a reflection layer alternately arranged are adhered on a plane of a photodiode array substrate to form a detection element array block.

Fig. 12 shows a conventional detector array. Focusing on one detector array block 90, the scintillator element 91 at the center of the block faces the X-ray tube focal point, while the scintillator element 92 at the end of the block faces the X-ray tube focal point. It will be. For this reason, the scintillator element 92 and the scattered ray removing collimator 93 are shifted from each other toward the end of the element array block, and the irradiation area seen from the focal point of the X-ray tube becomes wider. As a result, the amount of light emitted during X-ray irradiation increases, and the output changes. I do. In addition, since the path of the X-ray passing through the element differs between the scintillator element 91 at the center of the block and the scintillator element 92 at the end, the X-ray energy characteristics also change. For these reasons, variations occur in the light emission characteristics within the element array block, and artifacts tend to occur in the CT image.

 As a countermeasure against this, Japanese Patent Application Laid-Open No. 2000-237178 discloses that a plurality of detector element array blocks each composed of a plurality of elements are arranged in a chord shape by observing the circle around the focal point of the X-ray tube from the circumferential axis direction. This constitutes a channel detection element array. At this time, a solid-state detector is provided with a scattered radiation removing collimator in order to prevent scattered X-rays scattered by the subject from being incident on the detector element. A plurality of scattered radiation removing collimators are arranged in an arc shape so as to cover the gap between the detector elements and are accurately aligned with the detector element array. However, it is difficult to maintain accuracy when arranging a large number of detection element array locks in an arc shape. Furthermore, in order to orient each of the detection elements in the X-ray direction, the detector itself must be formed in an arc shape, but it is extremely difficult to create an arc shape detector. Disclosure of the invention

 In the radiation detector of the present invention, the angles of the plurality of reflective layers are particularly determined such that the plurality of phosphors fractionated by the plurality of reflective layers on the substrate are directed to the radiation source (focal point). .

That is, the radiation detector according to the present invention includes a substrate, a plurality of phosphor elements arranged on the substrate, and a reflection layer disposed between the adjacent phosphor elements to prevent crosstalk between the adjacent phosphor elements. A radiation detector comprising: a plurality of photoelectric conversion elements provided at positions corresponding to the phosphor elements on the substrate; and a contact surface between the phosphor element and the reflective layer; and a substrate at a central portion of the substrate. The angle between the substrate and the surface changes from the center of the substrate toward the end along the arrangement direction of the phosphor elements. According to this radiation detector, the orientation of the phosphor element defined by the phosphor element and the reflection layers on both sides of the phosphor element is substantially changed with respect to all the phosphor elements arranged in a plane. Since the incident direction can be the same, the apparent irradiation area of the radially incident radiation and the passage in the element can be made uniform, and the Variations in the output of the element can be eliminated.

 The angle formed between the contact surface between the phosphor element and the reflective layer and the substrate surface at the central portion of the substrate is, for example, with respect to the contact surface between the phosphor element and the reflective layer located at the central portion of the substrate or the position closest to the central portion. At 90 °, the distance from the radiation source to the top surface of the phosphor element is calculated from 90 ° for each phosphor array pitch, where Ld is the distance from the radiation source to the top of the phosphor element. Lp should be varied by the arrangement pitch of the phosphor elements.) Considering the processing accuracy of the radiation detector, an error of ± 0.2 ° can be tolerated.

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 center of the substrate (the center of the substrate) ) Is (η = 0, 1, 2, · · ·) in order from the center of the substrate, and 0 R _n = 90-a (n · θ) ± 0.2 ° (0 <a ≤ 2, η = 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 reflection layer located at the end of the substrate and the substrate surface at the center of the substrate (the angle at the center of the substrate). ) Is 0 s _n (η = 0, 1, 2 · · ·) in order from the center of the substrate, and 0 Sn = 90—a (n · θ) ± 0.2 ° (0 a a2, n = 0, l, 2 · · ·).

であった。 発明者らの検討によれば、 0 _Rnが 90° より小さければ本発明の効果があり、 また 9Ο_2η· 0 ° より小さければ、 逆に発 光特性が劣ってくることが分かった。 厳密に上記のような角度を有した素子ァレ ィを作製することは難しいため所定寸法誤差を許容範囲とした。 たとえば各要素 の傾斜角度が、 ±0.2° の精度で素子ァレイプロックを作製すれば充分な効果が得 られる。 0 for conventional detectors

Met. According to the studies by the inventors, it has been found that if 0 _Rn is smaller than 90 °, the effect of the present invention is obtained, and if 0 _Rn is smaller than 9Ο_2η · 0 °, on the contrary, the light emission characteristics are inferior. Since it is difficult to manufacture an element array having the above-described angle exactly, a predetermined dimensional error is set as an allowable range. For example, a sufficient effect can be obtained if an element array lock is manufactured with an inclination angle of each element of ± 0.2 °.

 In the radiation detector of the present invention, the width changes along the depth direction by changing the direction of the phosphor elements in the arrangement direction. This change in width can be provided in the phosphor element and / or the reflective layer. Hereinafter, the length in the direction in which the phosphor elements and the reflection layer are arranged is referred to as the width of the phosphor elements and the reflection layer, and the distance from the radiation incident surface to the substrate surface is referred to as the height or the depth.

A medical image diagnostic apparatus according to the present invention includes a radiation source, a radiation detector having the above-described features disposed to face the radiation source, and holding the radiation source and the radiation detector, and driving the radiation source around a subject. Rotating disk, and radiation detected by the radiation detector. Image reconstruction means for reconstructing an image of the tomographic image of the subject based on the intensity of the line.

 In the medical image diagnostic apparatus according to the present invention, variations in output in the array direction of the radiation detectors are suppressed, and therefore, artifacts due to variations in output, particularly due to variations in radiation energy dependence that are difficult to correct by software. Yes Artifacts can be effectively suppressed.

 According to the first feature of the present invention, a substrate having a plurality of photoelectric conversion elements arranged in at least one predetermined direction is disposed on the substrate so as to correspond one-to-one to the plurality of photoelectric conversion elements. A plurality of phosphor elements, and a reflective layer inserted between the phosphor elements arranged in the predetermined direction so as to prevent crosstalk between the phosphor elements. Provided is a radiation detector in which an angle between a contact surface between a phosphor element and the reflective layer and the substrate surface changes from the center to the end of the substrate along the arrangement direction.

 According to the second feature of the present invention, the angle between the contact surface between the phosphor element and the reflective layer and the substrate surface becomes smaller from the center to the end of the substrate.

 According to a third feature of the present invention, the width of the phosphor element in the predetermined direction is constant irrespective of the position in the depth direction of the phosphor element orthogonal to the predetermined direction. According to a fourth feature of the present invention, the width of the reflective layer in the predetermined direction is constant irrespective of the position in the depth direction of the reflective layer orthogonal to the predetermined direction.

 According to a fifth aspect of the present invention, the plurality of phosphor elements are formed of a single member, and the upper surface and the lower surface thereof are respectively aligned and parallel, and the reflection layer is formed of the plurality of phosphor elements. It is formed in a groove for partitioning the phosphor element.

 According to the sixth aspect of the present invention, the width of the groove for partitioning the plurality of phosphor elements is substantially constant irrespective of the position in the thickness direction, and the reflecting material is cured after filling with a curable material. Then, a metal plate or a metal plate was adhered in the groove.

 According to the seventh feature of the present invention, the width of the groove for partitioning the plurality of phosphor elements becomes thicker as the position in the thickness direction becomes deeper, and the reflecting material is cured after filling and shrinking a material that cures and contracts. did.

According to an eighth aspect of the present invention, the material to be cured is epoxy resin, acrylic resin It is at least one of inorganic resin powders such as titanium oxide, aluminum oxide, and barium sulfate contained in an organic resin such as lunar phenol resin.

 According to a ninth feature of the present invention, the material that cures and shrinks is an organic resin such as an epoxy resin, an acryl resin, or a phenol resin, and an inorganic compound powder such as titanium oxide, silicon oxide, or barium sulfate. It is at least one of those contained. According to a tenth feature of the present invention, a plurality of the above-described radiation detectors are prepared and juxtaposed in a one-dimensional or 2′-dimensional direction such that their long-axis directions intersect at a predetermined angle of 0 ° to 90 °. A radiation detector unit configured as follows.

 According to an eleventh aspect of the present invention, a radiation detector arranged to face a radiation source, a rotating disk that is driven to rotate while holding the radiation detector, and detected by the radiation detector 3.A medical image diagnostic apparatus comprising: image reconstruction means for reconstructing a tomographic image of an object based on the intensity of radiation, wherein the medical image diagnostic apparatus uses the radiation detector according to claim 2 as the radiation detector. I will provide a.

 According to a twelfth aspect of the present invention, a radiation detector arranged to face a radiation source, a rotating disk that is driven to rotate while holding the radiation detector, and detected by the radiation detector A medical image diagnostic apparatus, comprising: image reconstruction means for reconstructing a tomographic image of an object based on the intensity of radiation, wherein the medical image diagnosis uses the radiation detector unit according to claim 10 as the radiation detector. Provide equipment.

According to a thirteenth feature of the present invention, in the radiation detector of the medical image diagnostic apparatus, the shortest distance from the radiation source to the upper surface of the phosphor element is Ld, and the arrangement pitch of the phosphor elements in the predetermined direction is Ld. Where Lp and Θ = tan "i (Lp / Ld), and when the number of the phosphor elements in the predetermined direction is an even number, one of the reflective layers and the phosphor located on the end side of the substrate are disposed. If the angle between the contact surface between the element and the substrate surface and the substrate center side is 0 Rn (η = 0, 1, 2, ··-) in order from the substrate center, 0 _Rn = 9O— a ( n · θ) ± 0.2 ° (0 <a≤2η = 0,1,2, ·), and when the number of phosphor elements is odd, one of the above phosphor elements and its base The angle between the contact surface between the reflective layer located on the edge of the board and the surface of the board and the center of the board is 0 s _{n in} order from the center of the board. (η = 0,1,2 ···) 0 s _n = 9O— a (n · θ) ± 0.2 ° (0 <a≤2, η = 0,1,2 ···) An image diagnostic apparatus is provided. According to a fourteenth feature of the present invention, in each of the radiation detectors in the radiation detector unit of the medical image diagnostic apparatus, the shortest distance from the radiation source to the upper surface of the phosphor element is Ld, When the arrangement pitch of the elements in the predetermined direction is Lp, Θ = tan-i (Lp / Ld), when the number of the phosphor elements in the predetermined direction is even, one of the reflection layers and the end of the substrate are disposed. The angle between the contact surface between the phosphor element located on the side of the substrate and the surface of the substrate and the center of the substrate is 0 R _n (η = 0, 1, 2, · · ·) in order from the substrate center. Then 0 _Rn = 9O—a (n · θ) ± 0.2. (0 <a≤2 n = 0, l, 2 · · ·), and when the number of the phosphor elements is an odd number, one of the phosphor elements and the reflective layer Assuming that the angle between the contact surface between the substrate and the substrate surface at the center of the substrate is 0 s _n (n = 0,1,2-· ·) in order from the substrate center, 0 _Sn = 9O—a (η · Θ) ± 0.2. (0 <a≤2, n = 0, l, 2 ···). BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 is a cross-sectional view of a radiation detector according to one embodiment of the present invention. 2A and 2B are cross-sectional views of a main part of the radiation detection element array block in FIG. 1, wherein FIG. 2A shows a case where the number of detection elements is even, and FIG. 2B shows a case where the number of detection elements is odd. Figure 3 is a diagram showing the positional relationship between the radiation source and the radiation detection element array. FIG. 4 is a sectional view of a main part of a radiation detector according to another embodiment of the present invention. FIG. 5 is an example in which a plurality of detectors are arranged in the rotation axis direction according to the embodiment. FIG. 6 is a diagram illustrating an example of a method for manufacturing a radiation detector according to the present invention. FIG. 1 is a configuration diagram of an X-ray CT apparatus equipped with a radiation detector according to the present invention. FIG. 8 (a) is a diagram showing the X-ray energy characteristics of the radiation detector of the present invention, and FIG. 8 (b) is a diagram showing the X-ray energy characteristics of the conventional radiation detector. FIG. 9 shows a flow of manufacturing the detector according to the first embodiment of the present invention. FIG. 10 shows a flow of manufacturing a detector according to the second embodiment of the present invention. FIG. 11 shows a flow of manufacturing the detector according to the third embodiment. FIG. 12 is a diagram showing the relationship between a conventional radiation detecting element array and a scattered radiation removing collimator. BEST MODE FOR CARRYING OUT THE INVENTION Hereinafter, embodiments of the present invention will be described in detail with reference to the accompanying drawings.

 FIG. 1 is a cross-sectional view of a radiation detector according to one embodiment of the present invention. The radiation detector includes a substrate 11, a photodiode 14 (photoelectric conversion element) formed on the substrate, a plurality of scintillator elements 12 (phosphor elements), and a reflection layer 13. The scintillator elements 12 are arranged directly above the photodiodes 14 in a one-to-one relationship, and the scintillator elements 12 and the reflective layers 13 are alternately arranged in the width direction. Therefore, immediately below the reflective layer 13 is a portion of the substrate 11 where the photodiode 14 does not substantially exist. One scintillator element is constituted by the scintillator element 12 which is channel-separated by sandwiching both sides in the width direction with the reflective layer 13 and the photodiode 14 immediately below the scintillator element.

 The size of the scintillator element and the reflective layer and the number of scintillator elements arranged depend on the application of the radiation detector and are not particularly limited.For example, in the case of a detector array for an X-ray CT device, a substrate having a width of about 20 to 40 mm is used. The scintillator elements 12 are provided at an arrangement pitch Lp of about 1 mm. The width of the reflective layer 13 in the arrangement direction is about 1/10 to 210 of the normal arrangement pitch Lp.

 The radiation detector according to the present embodiment is configured so 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 almost perpendicular to the line connecting the scintillator element 12 and the radiation source (or the focal point, if it is an X-ray). To do. The arrangement of the scintillator element 12 and the reflective layer 13 of the radiation detector is bilaterally symmetric about the reflective layer 13 in FIG. 2A, and bilaterally symmetric about the scintillator element 120 in FIG. 2B.

When 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 Θno between the contact surface between the reflective layer 13 and the scintillator elements 121 on both sides and the substrate plane is 90 °. Scintillator elements 121, 121 of the angle theta _.kappa.1 the contact surface of the reflective layer and the scintillator elements in the end portion side on the outside makes with the substrate surface (angle of the substrate center side) as will be hereinafter both 90 ° I do. Thereafter, the angle between the contact surface and the substrate surface gradually decreases toward the end. I do.

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 0so between the contact surface of the scintillator element 120 and the reflective layers 131 on both sides with the substrate plane is 90 °. The angle 0 _S1 (the angle at the center of the substrate) formed by the contact surface between the scintillator elements 121 and 121 and the reflection layers 131 at the end further to be 90 ° or less. You. Thereafter, the angle between the contact surface and the substrate surface is gradually reduced toward the end.

 As described above, the angle is changed such 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 at the edge of the substrate and the substrate at the center of the substrate are described. When the angle between the plane and the plane (the angle at the center of the substrate) is (η = 0, 1, 2, · · ·) in order from the center of the substrate, Equation 1 is obtained.

 0 Rn-9O-a (η · Θ) ± 0.2. · · · (1)

 (Where 0 <a≤2, n = 0, l, 2 · ·, Θ = tan ~ i (Lp / Ld), Ld is the distance from the radiation source to the top of the scintillator element, and Lp is the The arrangement pitch is the same hereafter.)

If 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 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 defined as 0 _Sn (η = 0,1,2 · · ·) in order from the center of the substrate.

 0 Sn = 9O-a (η · ±) ± 0.2 ° (2)

Note that a in the above equations 1 and 2 is a coefficient. The center line of the scintillator points to the direction of the X-ray source when a = l with a value satisfying 0 <a≤2. By the way, a = 0 indicates the arrangement of scintillators according to the conventional technology, and the center lines of each scintillator point in the same direction in parallel. According to this embodiment, the effect starts when the center line of each scintillator is slightly inclined and 0 <a <l, and becomes ideal when a = l. <a≤2 is still effective.

 In Equations 1 and 2 above, Θ Otan-1 (Lp / Ld)) is equivalent to the elevation angle per pitch of the scintillator element, and thus the shape of the scintillator element for each elevation pitch By inclining the shape, the shape of the scintillator element always faces the radiation source. This can prevent radiation entering the adjacent scintillator element from entering, thereby suppressing variations in the amount of emitted light.

In Equations 1 and 2, ± 0.2 ° is a permissible error in consideration of dimensional errors in manufacturing. That is, according to the study of the present inventors, if an element array block is manufactured with an inclination angle of each element of ± 0.2 °, the effect of the present embodiment of the present invention can be obtained. If there is a variation, this effect is reduced. In addition, the effect of the present invention can be obtained when the inclination angle of each element is in the range of η < _0η ^ η · 0, but when the inclination angle 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 0 _n = 9O—a (n'θ) ± 0.2 ° (0 <a≤2, η = 0, 1, 2, ·).

The effects of the present embodiment will be described more specifically. For example, as shown in FIG. 3, a 24-channel detection element array block 32 having a scintillator element arrangement pitch Lp of 1 mm is arranged at a distance of Ld = 1037 mm from the focal point of the X-ray tube 31, that is, on an arc having a radius of 1037 mm. Suppose you have been. Here, when the detection element array block 32 is the conventional detection element array block 80, the reflective layer at the center of the substrate is configured to be parallel to the normal direction of the substrate. The reflective layer at the center of the substrate is oriented in the direction of the X-ray tube focal point, but the reflective layers located at positions other than the center of the substrate are 素 子 (= tan-1 (Lp / Ld)) = approx. ° deviates from the direction of the X-ray tube focal point. In other words, the reflection layer at the end of the block, which is 12 elements away from the center, is shifted by about 0.66 ° (= 12x0) from the direction of the X-ray focal point. Here, the angle between the normal to the side of the reflective layer at the center of the substrate (the surface that intersects the substrate surface) and the direction of the focal point of the X-ray tube is 90 degrees. Since the angle formed by the direction of the X-ray tube focal point (complementary angle in the case of an obtuse angle) is 0 R _n = 9O—η · 0, the normal to the side surface of the reflective layer at the end of the block and the X-ray tube focal point The angle between this direction and 90 ° is about 0.66 ° = about 89.34 °.

Naturally, the shape of the scintillator element is not Will be. The same applies to the case where an odd number of scintillator elements are arranged.The scintillator element in the center of the block has a shape facing the X-ray tube focal point, but the distance from the center of the substrate increases by approximately 0.055 °. The shape is shifted from the direction of the tube focus. Therefore, the angle between the normal to the side surface of the scintillator element and the direction of the X-ray tube focal point is 0 s _n = 9O_n · θ (η = 0, 1, 2, · · ·) in order from the center of the substrate.

 On the other hand, in the case of the detection element array block according to the present embodiment shown in FIG. 1, the angle formed between the side surface of the reflective layer and the substrate surface (the angle on the substrate center side) 0 Rn (Π =

0,1,2 ···) force S 0 _Rn = 9O— η · θ (η = 0,1,2 ···), the scintillator elements 12 and the reflective layer 13 are arranged, Even when the element array blocks are arranged on a polygonal shape (polygon) simulating an arc, all the elements can be arranged in the direction of the X-ray tube focal point.

 Note that FIGS. 1 and 2 show a case where the width of the reflective layer is constant at any position in the depth direction, but as shown in FIG. 4, the width of the reflective layer is changed in the depth direction. The width of the scintillator element may be constant at any position in the depth direction by changing the position. Although only one-dimensional array is shown in the figure, a two-dimensional element array can have a similar structure in each two-dimensional array direction.

 In the case of a two-dimensional element array, as shown in FIG. 5, a plurality of the above-described two-dimensional element arrays are rotated so that their major axes intersect at a predetermined angle of 0 ° or more and 90 ° or less. The two-dimensional element arrays can be arranged at different angles so that each two-dimensional element array faces the direction of the X-ray tube focal point. In this case, the variation in the detector array direction can be suppressed over a wider area, and the artifact caused by this can be reduced and eliminated.

 Next, as a method of manufacturing the radiation detecting apparatus according to the present embodiment, any of the following conventional methods (1) to (3) for manufacturing a channel-separated scintillator element block can be applied. . In particular, the method (3) is desirable in terms of ease of production and performance of the obtained element.

(1) A plurality of plate-shaped scintillator materials are laminated at predetermined Then, the laminated block is cut to a predetermined thickness and bonded to the substrate on which the photodiode array is formed.

 (2) A method in which a plate-shaped scintillator material is adhered to a photodiode array substrate, and then a cutout groove is formed at a predetermined pitch in the plate-shaped scintillator material, and a material constituting a reflective layer is placed in the groove.

 (3) Cut grooves are formed at a predetermined pitch in the plate-shaped scintillator material, and the material constituting the reflective layer is filled in this groove by a specific method. A method of bonding on a substrate on which is formed.

 Figure 6 illustrates the method (3). First, a notch groove 51 having a predetermined pitch, a predetermined width, and a predetermined depth is formed in a rectangular parallelepiped scintillator raw material 50 without leaving a bottom thereof. As the scintillator material, a phosphor having a high radiation-to-light conversion efficiency such as a rare-earth phosphor can be used. The phosphor is, for example, a rare-earth acid composed of a host crystal having a garnet structure containing at least Gd, Al, Ga, and O with Ce as a light-emitting element as described in Japanese Patent Application Publication No. JP-A-2001-004753. It is a scintillator of arsenal.

Next, the cut groove 51 is filled with a material 52 for forming a reflective layer. When the material 52 does not harden and shrink, each cut groove 51 is formed to have a predetermined angle with respect to the surface of the scintillator material. This angle is such that the angle formed by the contact surface between the reflective layer and the scintillator element and the substrate surface, when formed as a detection element array as described above, is inclined from the center to the end of the substrate. The processing can be achieved by sequentially changing the angle between the material and the cutting edge in a processing machine such as a multi-wire saw-slicer every time it approaches the end. On the other hand, when the material 52 is a material that cures and shrinks, the cut grooves 51 may be formed in parallel. In this case, the reflective layer tries to shrink due to curing shrinkage, but the dimensions of the bottom surface do not change because the bottom surface of the scintillator material is connected. Therefore, the reflective layer near the bottom of the scintillator does not shrink much, but shrinks greatly toward the top. As a result, a substantially wedge-shaped cross-sectional shape whose width changes in the depth direction as described above can be provided to the reflective layer. After the cut grooves 51 are formed in this way, a material 52 for forming a reflective layer is filled in the grooves. As the material 52 for forming the reflection layer, epoxy resin, acrylic resin, An organic resin such as an enol resin, a powder containing an inorganic compound powder such as titanium oxide, aluminum oxide, barium sulfate or the like as a reflective material, a metal plate having a high light reflectance, or a combination thereof. Can be used.

 Here, when a material containing a reflecting material in a resin is used as the material 52, it is centrifugally injected 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.

 Alternatively, in the case where a material in which a reflecting material is dispersed in an organic resin is used as the material 52, the reflecting layer is formed by hardening the organic resin in the material 52.

 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.

 When a material in which a reflective material is dispersed in an organic resin is used as the material 52, the resin shrinks when the organic resin hardens, but does not shrink because the bottom of the groove is fixed to the bottom of the scintillator material. Therefore, a slope is formed from the bottom to the top of the groove. 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 10000 20000 cPs is used as an organic resin, a groove having a width of 0.15 mm and a depth of 10 mm can form a slope of about 0.0400 °. 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 width in the depth direction as described above is substantially changed. A wedge-shaped cross-sectional shape can be imparted to the reflective layer.

 On the other hand, when a material without such curing shrinkage, for example, a metal plate, is used, a reflection layer having a desired angle can be formed by previously inclining the cut groove 51 itself.

 The radiation detector of the present invention is obtained by bonding the scintillator manufactured in this manner on the substrate 53 on which the photodiode array is formed in advance in the same arrangement pitch as the scintillator elements. A known photodiode such as a PIN photodiode can be used as the photodiode.

Next, an X-ray CT apparatus using the radiation detector of the present invention will be described. Figure 6 The outline of an X-ray CT apparatus is shown as an example of medical image diagnosis to which the present invention is applied. 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. X-ray tube 612 mounted on the X-ray tube, a collimator 613 mounted on the X-ray tube 612 to control the radiation direction of the X-ray flux, and an X-ray detector mounted on a rotating disk 611 facing the X-ray tube 612 615, 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. Have been.

 The image reconstruction unit 620 performs image processing for reconstructing a CT image by processing the measurement data S1 transmitted from the input device 621 and the detector circuit 616 for inputting the subject's name, examination date and time, examination conditions, and the like. Circuit 622, an image information adding unit 623 for adding information such as the subject name, examination date and time, and examination conditions input from the input device 621 to the CT image created by the image arithmetic circuit 622, and image information added. A display circuit 624 that adjusts the display gain of the CT image signal S2 and outputs the result to the display monitor 630.

 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. This X-ray is given directivity by a collimator 613 and 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.

The X-ray detector 615 is composed of a large number of detection element array blocks, each of which has a plurality of detection elements that combine a scintillator element (phosphor element) and a photodiode (photoelectric conversion element), arranged in a polygonal shape close to an arc. Consists of It is assumed that the radiation detector has a detection element array block according to the present embodiment as shown in FIG. 1, FIG. 2, or FIG. The radiation detector may be a one-dimensional detector in which elements are arranged in a row in the circumferential direction, or a two-dimensional detector in which elements are arranged in a row. In the case of a two-dimensional detector, the array direction parallel to the rotation axis of the rotating disk 611 (the body axis direction of the subject) Also, each element is arranged facing the direction of the X-ray tube focal point.

 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 configured to have a shape directed in the X-ray irradiation direction from the focal point because the scintillator element is divided by the scintillator element and the reflective layers on both sides thereof. Even in the case of a scintillator element, sufficient radiation is incident as in the case of the central scintillator element, and a uniform output can be obtained.

 Fig. 8 (a) shows the X-ray energy characteristics when the X-ray detector according to the embodiment of the present invention as shown in Fig. 1 is used, and Fig. 8 (b) shows the results using the conventional X-ray detector. The X-ray energy characteristics in each case are schematically shown. In the conventional X-ray detector shown in Fig. 8 (b), the X-ray energy characteristics in the element array block vary widely. This causes variations in the light emission characteristics, and artifacts are likely to occur in CT images. On the other hand, in the X-ray detector according to the present embodiment, the dispersion of the X-ray energy characteristics inside the block can be suppressed, and the dispersion of the emission characteristics can be reduced. For this reason, a high-quality CT image with few artifacts can be obtained.

 Hereinafter, embodiments of the radiation detector according to the present invention will be described.

Example 1>

 The manufacturing flow of the radiation detector according to the present embodiment will be described based on FIG.

_{Gds (Al, Ga) 5 0} 12: Ce ceramic scintillator from configured scintillator plate X- Y plane of the (X, Y and Z directions of dimensions each 26 X 30 X 1.8mm), 24 at lmm pitch The channel was attached to a substrate on which a one-dimensional photodiode array was formed. Then, using a 0.13 mm thick diamond grindstone, grooves parallel to the Y direction were machined at lmm pitch in the X direction of the scintillator plate. The depth of the groove at this time is 1.8 mm.

At this time, when machining the groove closest to one end of the scintillator plate, the setting was such that the grindstone entered at an angle of 0.66 ° with respect to the Z direction (the normal direction of the photodiode array substrate). Similarly, the grindstone was set to enter the groove at the angle of 0.605 ° in the second and subsequent grooves from the end. The angle of the grindstone was gradually reduced in steps of 0.055 ° toward the center. In this way, the center groove (13th groove) of the scintillator plate The angle of the grindstone becomes 0 °, the groove on the other end side gradually increases the angle of the grindstone in the opposite direction, and the grindstone becomes 0.66 ° in the groove closest to the other end (25th). The angle was set for machining.

 After the groove processing was performed with the inclination angle changed in this way, an epoxy resin mixed with T> 2 powder was poured into the groove and cured, forming a reflective layer and producing a detector element block (see Fig. 5). ). In this detector element block, since the scintillator material is adhered to the substrate and the groove serving as the reflective layer is processed with a grindstone having a constant thickness, the reflective layer has a shape in which both side surfaces are parallel as shown in FIG.

 Forty detector element blocks fabricated in this way were mounted on an X-ray CT system, and the emission characteristics (X-ray energy characteristics) and their variations were measured. The distance from the X-ray tube focal point to the detector element block was 1037 mm. The X-ray energy characteristics E were defined as follows by irradiating X-rays at a tube voltage of 100 kV and 120 kV and a tube current of 100 mA, and by the output ratio of each channel at that time.

 E = 6095 · log {(Rioo / Xioo) / (R120ZX120)}

_{Wherein, X 10 o (X120),} R100 (Ri2o) is sump lambda ^ when the respective tube voltages 100 kV (120 kV), which is the output of the reference.

 The variation in X-ray energy characteristics was defined as the difference between the maximum value and the minimum value of the characteristic の of each channel in the detector element block. If this variation is 5 or more, ring artifacts are likely to occur in CT images.

 The variation of the X-ray energy characteristics within the block of this example was about 0.3, where the variation of the conventional device block was 1. Therefore, it was shown that this X-ray CT device can suppress artifacts in CT images.

<Example 2>

The manufacturing flow of the radiation detector according to the present embodiment will be described with reference to FIG. _{Gd 3 (Al, Ga) 5} 0 12: Ce ceramic scintillator from configured scintillator plate onto the X-Y plane of the (X, Y and Z directions of dimensions each 24 X 36 X 4mm), diameter 0.13mm Using a multi-wire saw, a grid-like grooving process parallel to the X and Y directions was performed. In the Y direction, 23 grooves were machined in the X direction at lmm pitch, and in the X direction, 15 grooves were machined in the Y direction at 2.25mm pitch. Any depth of groove Was 3 mm, and it was connected at the bottom of the scintillator plate. Each groove was machined vertically to the bottom of the scintillator.

Thus the grooves of the scintillator plate subjected to groove processing, cured by pouring Epoki sheet resin mixed with Ti0 ₂ powder to form a reflective layer. At this time, curing and shrinkage of the epoxy resin occurred, but the scintillator elements were connected at the bottom, so that the shape became inclined toward the center. After that, the upper and lower surfaces were cut and polished to a thickness of 1.8 mm to produce a 24 × 16 channel scintillator element array. In this scintillator element array, the scintillator element was parallelized on both sides because it was burned by a wire saw, but the reflective layer had a shape in which the upper surface 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 ° for the end element in the X direction and 89.14 ° for the end element in the Y direction.

 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, if the variation in the conventional element block in both the X and Y directions is 1, the value is about 0.4, indicating that the X-ray CT system using this detector element block can suppress artifacts in CT images. .

 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 defined by a phosphor element (scintillator element) and reflection layers disposed on both sides thereof By providing the direction of the element substantially the same as the direction of incidence of radiation, it is possible to suppress variations in light emission characteristics within the detector block for radially incident radiation and provide a highly accurate radiation detector can do. Further, by providing such a radiation detector, it is possible to provide a medical image diagnostic apparatus such as an X-ray CT apparatus with less artifact.

<Example 3>

 A plurality of scintillator element arrays may be prepared, and the radiation detectors may be configured by juxtaposing them in a one-dimensional or two-dimensional direction such that their major axes intersect at a predetermined angle of 0 ° or more and 90 ° or less. .

A 24 x 8 channel scintillator element array was produced in the same process as in Example 2, A radiation detector was obtained by adhering to a 24 × 16 channel photodiode array substrate in a state where these two long axes were arranged in the channel direction at a predetermined angle. At this time, the photodiode of 24 × 16 channels was divided into regions of 24 × 8 channels, and each region was inclined at 0.50 ° with respect to the substrate surface.

 Even when this detector is used, as in the case of the second embodiment, each detector element can be a detection element array oriented in the direction of the X-ray tube focal point, and since it is juxtaposed at a predetermined angle, Variations in the detector array direction can be suppressed over a wider area, and artifacts due to this can be reduced and eliminated.

Claims

Claims: A substrate having a plurality of photoelectric conversion elements arranged in at least one predetermined direction, and a plurality of phosphor elements arranged on the substrate so as to correspond to the plurality of photoelectric conversion elements on a one-to-one basis. A reflection layer interposed between the phosphor elements arranged in the predetermined direction so as to prevent crosstalk between the phosphor elements and the reflection layer. A radiation detector wherein an angle between the substrate and the substrate surface changes from the central portion of the substrate toward the edge along the arrangement direction.  2. The radiation detector according to claim 1, wherein an angle between the contact surface between the phosphor element and the reflective layer and the substrate surface decreases from a central portion of the substrate toward an end portion. 3.  3. The radiation detector according to claim 2, wherein a width of the phosphor element in the predetermined direction is constant irrespective of a position in a depth direction of the phosphor element orthogonal to the predetermined direction.  The radiation detector according to claim 2, wherein a width of the reflection layer in the predetermined direction is constant irrespective of a position in a depth direction of the reflection layer orthogonal to the predetermined direction.  The plurality of phosphor elements are formed of a single member, and the upper surface and the lower surface thereof are respectively aligned and parallel, and the reflection layer is provided in a groove for dividing the member into the plurality of phosphor elements. 3. The radiation detector according to claim 2, wherein the radiation detector is formed.  The width of the groove for partitioning the plurality of phosphor elements is substantially constant irrespective of the position in the thickness direction, and the reflective material is formed by filling a material to be cured and then curing, or a metal plate is placed in the groove. 6. The radiation detector according to claim 5, wherein the radiation detector is bonded. The width of the groove for partitioning the plurality of phosphor elements becomes thicker as the position in the thickness direction becomes deeper, and the reflecting material is obtained by filling and curing a material that cures and contracts. 3. A radiation detector according to claim 1. The material to be cured is a mixture of an organic resin such as an epoxy resin, an acrylic resin, and a phenol resin, and a powder of an inorganic compound such as titanium oxide, aluminum oxide, and barium sulfate. The radiation detector according to 6.  The material to be cured is a material in which an inorganic resin powder such as titanium oxide, aluminum aluminum oxide and barium sulfate is contained in an organic resin such as an epoxy resin, an acrylic resin, and a phenol resin. A radiation detector according to item 7.  A radiation detector comprising a plurality of radiation detectors according to claim 2, which are arranged side by side in a one-dimensional or two-dimensional direction such that their major axes intersect at a predetermined angle of 0 ° or more and 90 ° or less. Instrument unit.  A radiation detector arranged to face the radiation source, a rotating disk holding and rotating the radiation detector, and a tomographic image of the object based on the intensity of the radiation detected by the radiation detector 3. A medical image diagnostic apparatus comprising: an image reconstructing means for reconstructing an image of the medical image, wherein the radiation detector according to claim 2 is used as the radiation detector. A radiation detector arranged to face the radiation source, a rotating disk holding and rotating the radiation detector, and a tomographic image of the object based on the intensity of the radiation detected by the radiation detector A medical image diagnostic apparatus comprising: an image reconstructing means for reconstructing an image of the medical image, wherein the radiation detector unit according to claim 10 is used as the radiation detector. The radiation detector of the medical image diagnostic apparatus according to claim 11, wherein a shortest distance from the radiation source to the upper surface of the phosphor element is Ld, an arrangement pitch of the phosphor elements in the predetermined direction is Lp, and Θ = tan. -i (Lp / Ld), when the number of the phosphor elements arranged in the predetermined direction is an even number, the contact between one of the reflective layers and the phosphor element located on the edge side of the substrate is determined. Assuming that the angle between the contact surface and the surface of the substrate at the center of the substrate is 0 Rn (η = 0, 1, 2, · · ·) in order from the center of the substrate, 0 R _n = 9O—a (n · Θ) ± 0.2 ° (0 <a≤2, η = 0,1,2 · · ·) and when the number of phosphor elements is odd The angle of the center of the substrate between the contact surface between one of the phosphor elements and the reflective layer located on the side of the substrate and the surface of the substrate is 0s _n ( If η = 0, 1, 2, ···), then 0 _Sn = 9O—a (n · θ) ± 0.2. (0 <a≤2, η = 0,1,2,...). 13. In each of the radiation detectors in the radiation detector unit of the medical image diagnostic apparatus according to claim 12, the shortest distance from the radiation source to the upper surface of the phosphor element is Ld, and the arrangement of the phosphor elements in the predetermined direction. When the pitch is Lp and Θ = tan-i (Lp / Ld), when the number of the phosphor elements arranged in the predetermined direction is an even number, the phosphor element is located on one of the reflection layers and the end of the substrate. Assuming that the angle between the contact surface between the phosphor element and the substrate surface on the substrate center side is 0R _D (η = 0, 1, 2, · · ·) from the substrate center, 0 _Rn = 90- a (n · θ) ± 0.2. (0 <a≤2η = 0, 1, 2,...), And when the number of the phosphor elements is odd, one of the phosphor elements and the reflection located on the end side of the substrate are disposed. If the angle between the contact surface between the layer and the substrate surface at the center of the substrate is 0s _n (η = 0, 1,2 · · ·) in order from the center of the substrate, then 0s _n = 9O—a (η · Θ) ± 0.2 ° (0 <a≤2, n = 0, l, 2 · ·-).