US20060039528A1 - Light detector, radiation detector and radiation tomography apparatus - Google Patents
Light detector, radiation detector and radiation tomography apparatus Download PDFInfo
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- US20060039528A1 US20060039528A1 US10/920,464 US92046404A US2006039528A1 US 20060039528 A1 US20060039528 A1 US 20060039528A1 US 92046404 A US92046404 A US 92046404A US 2006039528 A1 US2006039528 A1 US 2006039528A1
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
- G01T1/164—Scintigraphy
- G01T1/1641—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras
- G01T1/1644—Static instruments for imaging the distribution of radioactivity in one or two dimensions using one or several scintillating elements; Radio-isotope cameras using an array of optically separate scintillation elements permitting direct location of scintillations
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
- A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
- A61B6/02—Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
- A61B6/03—Computerised tomographs
- A61B6/032—Transmission computed tomography [CT]
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- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/29—Measurement performed on radiation beams, e.g. position or section of the beam; Measurement of spatial distribution of radiation
- G01T1/2914—Measurement of spatial distribution of radiation
- G01T1/2985—In depth localisation, e.g. using positron emitters; Tomographic imaging (longitudinal and transverse section imaging; apparatus for radiation diagnosis sequentially in different planes, steroscopic radiation diagnosis)
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Abstract
Description
- The present invention relates to a light detector, a radiation detector and a radiation tomography apparatus.
- An X-ray detector wherein X-ray detection modules each formed with a plurality of photodiodes are provided side by side in plural form in a channel direction, has been used in a multi-slice type X-ray CT apparatus (refer to, for example, the following patent document 1).
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FIG. 7 is a fragmentary plan view of a conventional X-ray detector. As shown inFIG. 7 , a plurality ofphotodiodes 243 are disposed in matrix form in the X-ray detector, more specifically, one X-ray detection module of the X-ray detector. In the drawing, a z direction corresponds to a slice direction (body axial direction) and an x direction corresponds to a channel direction, respectively. - Since wirings (signal lines) 245 for fetching signal charges generated in the
photodiodes 243 hardly pass light, they have heretofore been disposed in regions among thephotodiodes 243 respectively. - [Patent Document 1] Japanese Unexamined Patent Publication No. 2000-60840
- However, the multi-slice type X-ray CT apparatus is accompanied by the problem that as the number of the
photodiodes 243 in the slice direction (z direction) increases, the number of thewirings 245 that pass through the regions among the photodiodes arranged in the x direction increases, and there is hence a need to form fine wirings in the regions among thephotodiodes 243, thereby causing a difficulty in its fabrication and a rise in its manufacturing cost. - Here, it is also considered that in view of a limit of wiring miniaturization or scale-down, as shown in
FIG. 8 , the interval d defined betweenadjacent photodiodes 243 arranged in an x direction is made wide and the width w of eachphotodiode 243 is made small. A problem, however, arises in that since the area of thephotodiode 243 becomes small correspondingly, light detection efficiency is reduced. - Therefore, a first object of the present invention is to provide a light detector capable of suppressing a decrease in the area of a light receiving section with an increase in the number of wirings and thereby suppressing a reduction in light detection efficiency.
- A second object of the present invention is to provide a radiation detector capable of suppressing a decrease in the area of a light receiving section with an increase in the number of wirings and thereby suppressing a reduction in radiation detection efficiency.
- A third object of the present invention is to provide a radiation tomography apparatus capable of suppressing a decrease in the area of a light receiving section with an increase in the number of wirings and thereby suppressing a reduction in radiation detection efficiency.
- In order to achieve the above objects, a light detector of the present invention comprises a plurality of light receiving sections which are formed in a substrate and generate signal charges corresponding to the amount of incident light, and a plurality of wirings which are formed on the substrate and fetch the signal charges from the light receiving sections. Some of the plurality of wirings are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges.
- In the light detector of the present invention, some of the plurality of wirings are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges. The area of each light receiving section formed in the substrate is not limited by the wirings. Although the wirings shields the light from entering the light receiving sections, the light reaches the light receiving sections in regions other than the wirings, and thereby the signal charges corresponding to the amount of the incident light are generated by the light receiving units.
- In order to achieve the above objects, a radiation detector of the present invention comprises a plurality of light receiving sections which are formed in a substrate and generate signal charges corresponding to the amount of incident light, a plurality of wirings which are formed on the substrate and fetch the signal charges from the light receiving sections, and scintillators which are provided on the light receiving sections of the substrate and emit lights each having a wavelength longer than that of radiation in accordance with the incidence of the radiation. Some of the plurality of wirings are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges.
- In the radiation detector of the present invention, some of the plurality of wirings are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges. The area of each light receiving section formed in the substrate is not limited by the wirings. Although the wirings shields the lights emitted from the scintillators in accordance with the incidence of the radiation from entering the light receiving sections, the lights reach the light receiving sections in regions other than the wirings and thereby the signal charges corresponding to the amount of the incident light are generated by the light receiving sections.
- In order to achieve the above objects, a radiation tomography apparatus of the present invention comprises radiation irradiating means which irradiates a subject with radiation, and a radiation detector which detects the radiation transmitted through the subject. The radiation detector includes a plurality of light receiving sections which are formed in a substrate and generate signal charges corresponding to the amount of incident light, a plurality of wirings which are formed on the substrate and fetch the signal charges from the light receiving sections, and scintillators which are provided on the light receiving sections of the substrate and emit lights each having a wavelength longer than that of the radiation in accordance with the incidence of the radiation. Some of the plurality of wirings are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges.
- In the radiation tomography apparatus of the present invention, some of the plurality of wirings connected to the light receiving sections of the radiation detector are disposed so as to overlap with other light receiving sections different from the light receiving sections connected to fetch the signal charges, and the area of each light receiving section formed in the substrate is not limited by the wirings. The radiation is launched into the subject by the radiation irradiating means. The radiation transmitted through the subject is launched into the scintillators of the radiation detector. Although the wirings shields the lights emitted from the scintillators in accordance with the incidence of the radiation from entering into the light receiving sections, the lights reach the light receiving sections in regions other than the wirings, and thereby the signal charges corresponding to the amount of the incident light are generated by the light receiving sections.
- According to the light detector of the present invention, it is possible to suppress a decrease in the area of each of light receiving sections with an increase in the number of wirings and thereby suppress a reduction in light detection efficiency. According to the radiation detector of the present invention, it is possible to suppress a decrease in the area of each of light receiving sections with an increase in the number of wirings and thereby suppress a reduction in radiation detection efficiency. According to the radiation tomography apparatus of the present invention, it is possible to suppress a decrease in the area of each of light receiving sections with an increase in the number of wirings and thereby suppress a reduction in radiation detection efficiency. Further objects and advantages of the present invention will be apparent from the following description of the preferred embodiments of the invention as illustrated in the accompanying drawings.
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FIG. 1 is a schematic configurational view of a radiation tomography apparatus according to the present embodiment. -
FIG. 2 is a view showing detailed configurations of an X-ray tube and an X-ray detector. -
FIG. 3 (a) is a cross-sectional view of one X-ray detection module as view in a y-z plane, andFIG. 3 (b) is a plan view of the X-ray detection module, respectively. -
FIG. 4 is a plan view for describing the layout of photodiodes and wirings formed in a substrate. -
FIG. 5 is a cross-sectional view corresponding to line A-A′ ofFIG. 4 . -
FIG. 6 (a) is a fragmentary plan view for describing connections of wirings and photodiodes, andFIG. 6 (b) is a cross-sectional view taken along line B-B′ ofFIG. 6 (a), respectively. -
FIG. 7 is a fragmentary plan view of a conventional X-ray detector. -
FIG. 8 is a fragmentary plan view for describing problems of the conventional X-ray detector with an increase in the number of wirings. - Embodiments of the present invention will hereinafter be described with reference to the accompanying drawings.
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FIG. 1 is a schematic configurational view of a radiation tomography apparatus (X-ray CT apparatus) according to an embodiment. The X-rayCT apparatus 100 according to the present embodiment is equipped with anoperating console 1, an imaging table 10 and ascanning gantry 20. - The
operating console 1 is provided with aninput device 2 which receives an input from an operator, acentral processing unit 3 which executes an image reconstructing process or the like, a data acquisition buffer 5 which collects projection data obtained by thescanning gantry 20, aCRT 6 which displays a CT image reconstructed from the projection data, and amemory device 7 which stores programs, data and an X-ray CT image therein. - The imaging table 10 includes a
cradle 12 which carries a subject placed thereon in a bore (cavity portion) of thescanning gantry 20 and carries out it therefrom. Thecradle 12 is elevated by a motor built in the imaging table 10 and moves linearly along the table. - The
scanning gantry 20 is provided with an X-ray tube (X-ray irradiating means) 21, anX-ray controller 22, acollimator 23, an X-ray detector (radiation detector) 24, a data acquisition system (DAS) 25, arotating section controller 26 which rotates theX-ray tube 21 or the like about a body axis of the subject, and acontrol controller 29 which performs a transfer of a control signal or the like between theoperating console 1 and the imaging table 10. - A configuration of the X-ray CT apparatus according to the present embodiment is generally as described above. In the X-ray CT apparatus having the above configuration, the collection of projection data is performed in the following manner, for example.
- The position of the subject as viewed in a z-axis direction is fixed in a state in which the subject is placed in the cavity portion of a rotating
section 15 of thescanning gantry 20. An X-ray beam from theX-ray tube 21 is applied to the subject (projection of X-rays), and the X-rays transmitted through the subject are detected by theX-ray detector 24. Then, the detection of the transmitted X-rays is performed such that data corresponding to 360° are collected in directions of plural N (e.g., N=1,000) views while theX-ray tube 21 and theX-ray detector 24 are being rotated about the subject (i.e., a projection angle (view angle) is being changed). - The detected respective transmitted X-rays are converted into digital values by the DAS (Data Acquisition System) 25, which in turn are transferred to the
operating console 1 via the data acquisition buffer 5 as projection data. This operation is called “one scan”. A scan position is sequentially moved a predetermined amount in the z-axis direction (slice direction and body axial direction) and the next scan is performed. Such a scan system is called a conventional scan system (or axial scan system). However, a system for moving the imaging table 10 at a predetermined speed in synchronization with a change in the projection angle and collecting projection data while a scan position is being moved (theX-ray tube 21 and theX-ray detector 24 helically orbit around the subject), is referred to as a so-called helical scan system. The present invention can be applied even to both of the conventional scan system and the helical scan system. - The
operating console 1 stores the projection data transferred from thescanning gantry 20 in thememory device 7. Further, the operatingconsole 1 performs, for example, a predetermined reconstruction function and a superposition arithmetic operation and thereby reconstructs a tomographic image according to a back projection process. Here, the operatingconsole 1 is capable of reconstructing a tomographic image in real time on the basis of the projection data sequentially transferred from thescanning gantry 20 during scan processing and always displaying the latest tomographic image on theCRT 6. Further, the operatingconsole 1 invokes the projection data stored in thememory device 7 to thereby enable an image reconstruction anew. -
FIG. 2 is a diagram showing detained configurations of theX-ray tube 21 and theX-ray detector 24. - As shown in
FIG. 2 , theX-ray detector 24 is configured in such a manner that a plurality ofX-ray detection modules 240 are arranged on a circular arc with theX-ray tube 21 as the center. As described above, theX-ray tube 21 and theX-ray detector 24 rotate around the subject within an x-y plane, for example. In the specification of the present application, the positional relationship between theX-ray detection modules 240 at the central portion of theX-ray detector 24 is used to refer to an arcuate direction, i.e., an x direction of theX-ray detector 24 as a channel direction. Incidentally, the z direction corresponds to the slice direction (body axial direction). -
FIG. 3 (a) is a cross-sectional view of oneX-ray detection module 240 as viewed in a y-z plane, andFIG. 3 (b) is a plan view of theX-ray detection module 240, respectively. - As shown in
FIG. 3 , theX-ray detection module 240 is configured in such a manner that asubstrate 242 made up of silicon or the like is attached onto a central portion of acircuit board 241 comprising a ceramic board formed with wirings, for example. - The
substrate 242 is formed withphotodiodes 243 in matrix form. Each of thephotodiodes 243 comprises a p type impurity region formed in an n-type substrate 242, for example. A signal charge corresponding to the amount of incident light is generated within the substrate and captured by the correspondingphotodiode 243. -
Scintillators 246 fixed to thesubstrate 242 with an unillustrated transparent adhesive interposed therebetween are provided over thesubstrate 242 so as to correspond to thephotodiodes 243 respectively. Each of thescintillators 246 comprises a fluorescent material which reacts with an incident X-ray and thereby generates light having a wavelength longer than that of an X-ray, i.e., light in a substantially visible region which enables the generation of a signal charge by the photodiode. - A
reflection layer 247 is formed so as to cover thescintillators 246 on the X-ray incident side from above and between thescintillators 246. Thereflection layer 247 is made up of, for example, TiO2 which causes the X-rays to pass therethrough and reflects lights emitted from thescintillators 246. - Interconnections or wirings to be described later formed in the
substrate 242 are drawn out to both ends of thesubstrate 242 and connected to their corresponding circuits of thecircuit board 241 at both ends by wires. -
FIG. 4 is a plan view for describing the layout of thephotodiodes 243 andwirings 245 formed in thesubstrate 242. - As shown in
FIG. 4 , thewirings 245 made of aluminum or the like are connected to theircorresponding photodiodes 243 to fetch signal charges from thephotodiodes 243 arranged in matrix form. Thewirings 245 are formed between thephotodiodes 243 and thescintillators 246 respectively. Although only one side, i.e., the left side of a boundary line M at the central portion of a matrix of thephotodiodes 243 is illustrated inFIG. 4 ,photodiodes 243 are arranged even on the right side in a manner similar to it. - On the one side (left side) of the boundary line M, the
wirings 245 respectively connected to thephotodiodes 243, which have been arranged in matrix form, are drawn to the left end (one end) of thesubstrate 242. Thewirings 245 are roughly divided intowirings 245 a disposed between thephotodiodes 243 andwirings 245 b disposed so as to overlap withother photodiodes 243. Incidentally, particularly when it is not necessary to distinguish between thewirings 245 a and thewirings 245 b, they are simply called thewirings 245. - In the present embodiment, the wirings connected to the
photodiodes 243 on the center side (side close to the boundary line M) of the matrix extend in a z direction among thephotodiodes 243 arranged in a column direction (x direction) and are drawn to the end of thesubstrate 242. - The
wirings 245 b connected to thephotodiodes 243 on the end side of the matrix extend in the z direction so as to overlap with thephotodiodes 243 adjacent to one another in the z direction and are drawn to the end of thesubstrate 242. - Incidentally, although not shown in the drawing,
wirings 245 respectively connected to thephotodiodes 243 are drawn to the other end (right end) of the substrate in like manner even on the other side, i.e., right side of the boundary line M at the central portion of the matrix. -
FIG. 5 is a cross-sectional view corresponding to line A-A′ ofFIG. 4 . - As shown in
FIG. 5 ,photodiodes 243 each comprising a p-type impurity region are formed in asubstrate 242 on acircuit board 241, which is made of n-type silicon, for example. Incidentally, described more specifically, a region with a pn junction as the center serves as a photodiode. - An insulating
film 244 made of, for example, silicon oxide or the like is formed over thephotodiodes 243, and wirings 245 are formed over the insulatingfilm 244. The insulatingfilm 244 holds insulation between thephotodiodes 243 other than those intended for connection and thewirings 245. -
Scintillators 246 are fixed onto thesubstrate 242 formed with thewirings 245 with atransparent adhesive 248 interposed therebetween, and areflection layer 247 is formed so as to cover thescintillators 246. - As shown in
FIG. 5 , somewirings 245 b of thewirings 245, which are respectively located between thephotodiodes 243 and thescintillators 246, are formed over thephotodiodes 243, whereas theother wirings 245 a thereof are formed over a region between theadjacent photodiodes -
FIG. 6 (a) is a fragmentary plan view for describing connections between wirings and photodiodes, andFIG. 6 (b) is a cross-sectional view taken along line B-B′ ofFIG. 6 (a), respectively. - As shown in
FIG. 6 , connecting holes are defined in an insulatingfilm 244 lying overphotodiodes 243, and a wiring material such as aluminum is embedded into each of the connecting holes, whereby connectingportions 245 c for connectingwirings 245 and theircorresponding photodiodes 243 are formed. - The connecting
portions 245 c may be configured integrally with thewirings 245 b made of aluminum or the like. Alternatively, the connectingportions 245 c may be constituted of a material different from the wiring material so as to be embedded into the connecting holes. Incidentally, the connections ofwirings 245 a and theircorresponding photodiodes 243 are also similar to the above. - The operation of the
X-ray detector 24 will be explained. - X-rays, which pass through a subject and are thereby decayed, are launched into their
corresponding scintillators 246, so lights are emitted from thescintillators 246. The lights emitted from thescintillators 246 enter thephotodiodes 243. - When the lights enter into the
photodiodes 243, signal charges are produced, which in turn are captured by thephotodiodes 243. The signal charges captured by thephotodiodes 243 are fetched out to the end of thesubstrate 242 through thewirings 245, followed by being transmitted to a detection circuit of thecircuit board 241 via the wires. - Since the
X-ray detection modules 240 are arranged in large numbers adjacent to one another in the channel direction (x direction) as shown inFIG. 2 , the signal charges are taken out every channels in the z direction (slice direction) different from their arrangement or layout direction (x direction). - Since the light input is smaller because of structure of
scintillator 246 andreflector 247 and also the probability that the signal charge will reach the region of eachphotodiode 243 is low even if the signal charge occurs, in the region between thephotodiode 243 and thephotodiode 243 upon the operation of theX-ray detector 24, light is not detected very efficiently. Thus, the area of thephotodiode 243 may preferably be wide to capture the signal charge produced in thephotodiode 243 with efficiency. - When it is necessary to form the number of wirings larger than the number of wirings reasonably formable in the regions among the
photodiodes 243, the wirings are caused to lead so as to overlap with theperipheral photodiodes 243 and configured so as to take out the signal charges in the z direction (slice direction). - Thus, since the
photodiodes 243 exist among the wirings, although the light is cut off by the wirings per se, the detection of light at their portions is ensured. Therefore, fine or narrow wirings are no longer used at random and a reduction in light detection efficiency can be suppressed to the minimum. - The
reflection layer 247 is formed so as to cover thescintillators 246. Therefore, if the light is reflected by each of thewirings 245 b on thephotodiodes 243, the light is reflected by thereflection layer 247 and enters eachphotodiode 243. Thus, it is possible to prevent a reduction in light detection efficiency due to the existence of thewirings 245 b on thephotodiodes 243. - As shown in
FIG. 4 , the wirings connected to thephotodiodes 243 on the center side (side close to the boundary line M) of the matrix are placed so as to extend in the z direction among thephotodiodes 243 by priority. Further, thewirings 245 b connected to thephotodiodes 243 on the end side of the matrix are disposed in the form extended in the z direction so as to overlap with thephotodiodes 243 adjacent to one another in the z direction, whereby the number of thephotodiodes 243 on which thewirings 245 b are superimposed, can be suppressed as much as possible. Therefore, it is possible to suppress a reduction in light detection efficiency due to thewirings 245 b. - According to a radiation detector according to the present embodiment, as described above, it is possible to suppress a reduction in the area of a light receiving section with an increase in the number of wirings and suppress a decrease in radiation detection efficiency. Thus, according to a radiation tomography apparatus that adopts the radiation detector according to the present embodiment, it is possible to suppress a reduction in the area of a light receiving section with an increase in the number of wirings and suppress a decrease in radiation detection efficiency.
- The present invention is not limited to the description of the present embodiment. Although the present embodiment has explained the radiation detector and the radiation tomography apparatus, the present invention is applicable even to a light detector free of scintillators. Even in this case, it is possible to suppress a decrease in the area of a light receiving section with an increase in the number of wirings and suppress a reduction in light detection efficiency. Numerical values and materials mentioned in the present embodiment are illustrated by way of example. They are not necessarily limited to the illustrated ones. Many widely different embodiments of the invention may be configured without departing from the spirit and the scope of the present invention. It should be understood that the present invention is not limited to the specific embodiments described in the specification, except as defined in the appended claims.
Claims (16)
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JP5475686B2 (en) * | 2008-01-15 | 2014-04-16 | コーニンクレッカ フィリップス エヌ ヴェ | Solid state radiation detector |
US8610079B2 (en) * | 2009-12-28 | 2013-12-17 | General Electric Company | Robust radiation detector and method of forming the same |
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