JP4691074B2 - Radiation detection system - Google Patents

Description

  The present invention relates to a radiation detector, a radiation detection system, or an X-ray computer tomography apparatus (hereinafter referred to as an X-ray CT apparatus) including the radiation detector.

  The X-ray CT apparatus has an X-ray tube and a radiation detector. X-rays generated by the X-ray tube pass through the subject and enter the radiation detector. The detector includes a plurality of detection elements that detect X-rays as electrical signals. The detection element includes a phosphor such as a scintillator that converts X-rays into light, and a photoelectric conversion element such as a photodiode that converts the light into electric charges (electric signals). Recently, it has been studied to employ a semiconductor element that directly converts X-rays into electric charges as a detection element.

  In recent years, multi-slice radiation detectors have appeared. The multi-slice radiation detector includes a plurality of detector element arrays arranged in parallel along the slice direction. Each of the detection element arrays has a plurality of detection elements arranged in a line in the channel direction substantially orthogonal to the slice direction.

  This multi-slice type radiation detector is required to increase the number of rows. However, the conventional radiation detector has a limited number of rows.

  The biggest factor that hinders the increase in the number of rows of radiation detectors is the wiring structure and the connection structure. Here, for convenience of explanation, it is assumed that the photodiodes are arranged in a matrix of n × m (channel direction × slice direction). That is, n photodiodes are arranged in the channel direction, and m photodiode arrays are arranged in parallel in the slice direction.

  The plurality of photodiodes and the plurality of switching elements are connected via a plurality of signal lead lines. M signal lead lines of m photodiodes arranged in the slice direction are formed in a gap between adjacent photodiodes in the channel direction.

  Therefore, the number of photodiode columns is determined depending on the number of signal lead lines that can be formed in the gap between adjacent photodiodes in the channel direction. If the gap is enlarged, the number of signal lead lines can be increased.However, in this case, the area of the photodiode sensitive area is reduced in inverse proportion to the gap enlargement, so the sensitivity decreases. End up.

  An object of the present invention is to provide a radiation detector, a radiation detection system, and an X-ray CT apparatus that can increase the number of rows.

An aspect of the present invention includes a scintillator that converts X-rays incident from the surface side into light, at least one photodiode chip that includes a plurality of photodiodes that convert the converted light into electrical signals, and the plurality of the plurality of photodiodes. And at least one switching chip having a plurality of switching elements for reading a plurality of signals from the photodiode, and at least one data having a plurality of data collecting sections for amplifying and digitizing the read signals. A rigid multilayer wiring board that commonly mounts the collection chip, the photodiode chip, the switching chip, and the data collection chip, and the back surface of the switching chip is coupled to the surface of the multilayer wiring board; The back surface of the photodiode chip is coupled to the surface of the switching chip. Is, the data acquisition chips on the back surface of the multilayer wiring board is characterized in that it is coupled.

  According to the present invention, the number of columns can be easily increased.

  Hereinafter, embodiments of the present invention will be described with reference to the drawings. The present embodiment relates to a two-dimensional array type radiation detector and an X-ray CT apparatus (X-ray computed tomography apparatus) equipped with the radiation detector. In the X-ray CT apparatus, a rotation / rotation (ROTATE / ROTATE) type in which an X-ray tube and a radiation detector are rotated as one body, and a number of detection elements are arrayed in a ring shape. There are various types such as a fixed / rotation (STATIONARY / ROTATE) type in which only the tube rotates around the subject, and the present invention can be applied to any type. Here, the rotation / rotation type that currently occupies the mainstream will be described.

  In addition, in order to reconstruct one volume of voxel data (which constitutes one volume data) (or one tomographic image) (both will be described later), a projection of about 360 ° around the subject is performed. The projection data corresponding to about 210 to 240 ° is also required for the data even in the half scan method. The present invention can be applied to any method. Here, a description will be given assuming that one volume of voxel data (or one tomographic image) is reconstructed from the projection data of about 360 ° of the general former.

  In addition, the mechanism for converting incident X-rays into electric charge includes an indirect conversion type in which X-rays are converted into light by a phosphor such as a scintillator and the light is converted into electric charge by a photoelectric conversion element such as a photodiode, and a specific semiconductor The mainstream is the generation of electron-hole pairs in semiconductors by the X-rays and the transfer to the electrodes, that is, the direct conversion type utilizing the photoconductive phenomenon. Any of these methods may be employed as the X-ray detection element, but here, the former indirect conversion type will be described.

  In the following, the width of the sensitive area of the photodiode is defined as a converted value on the rotation center axis of the X-ray tube. That is, “a photodiode having a sensitive area width of 1 mm” means “a photodiode having a sensitive area width corresponding to 1 mm on the rotation center axis of the X-ray tube”, and X-rays diffuse radially. The width of the actual sensitive area of the photodiode is less than 1 mm according to the ratio of the actual distance between the X-ray focal point and the photosensitive area of the photodiode with respect to the distance between the X-ray focal point and the rotation center axis. Slightly wider.

(First embodiment)
FIG. 1 is a perspective view of the radiation detector according to the first embodiment. The radiation detector is composed of a plurality of detector modules 1 arranged in a substantially arc shape along the slice direction. FIG. 2A schematically shows the configuration of one detector module 1. FIG. 2B shows a cross section of one detector module 1.

  Each detector module 1 includes a scintillator 9 that converts incident radiation (here, X-rays) into light, at least one photodiode chip 2 that converts the converted light into an electrical signal, and an electrical signal from the photodiode chip 2. At least one switching chip 4 for reading out and at least one DAS chip (data acquisition system chip) 8 for amplifying and digitizing the read-out electric signal, a single rigid printed wiring board made of ceramic 10 in common. The rigid printed wiring board 10 is long in the slicing direction, and the photodiode chip 2, the switching chip 4, and the DAS chip 8 are arranged on the rigid printed wiring board 10 along the slicing direction.

  The photodiode chip 2 includes a plurality of photodiodes 3 and a plurality of Al wirings 14 formed on the surface of a silicon substrate. N photodiodes 3 are arranged with a certain gap in the slice direction. With respect to the channel direction, M photodiodes 3 are arranged with the same constant gap. Note that the slice direction is substantially parallel to the body axis direction of the subject. The channel direction is substantially orthogonal to the body axis direction of the subject.

  The plurality of Al wirings 14 are respectively connected to the plurality of photodiodes 3. The (N / 2) × M photodiodes 3 arranged in the right half area of the photodiode chip 2 with respect to the slice direction are formed in a gap substantially parallel to the slice direction (N / 2) × M. It is pulled out to the right outside through the Al wiring 14. (N / 2) Al wirings 14 corresponding to (N / 2) photodiodes 3 in the same slice row are formed in a high density in a gap between adjacent slice rows.

  The (N / 2) × M photodiodes 3 arranged in the left half area of the photodiode chip 2 with respect to the slice direction are (N / 2) × M formed in a gap substantially parallel to the slice direction. It is pulled out to the left outside through the Al wiring 14. (N / 2) Al wirings 14 corresponding to (N / 2) photodiodes 3 in the same slice row are formed in a high density in a gap between adjacent slice rows.

  The Al wirings 14 drawn to the left and right are connected in common to the signal readout lines 7 by (N / 2) lines through the transistors 5 respectively.

  The number of wirings 14 that can be formed in the gap is determined by the width of the gap, the thickness of the wirings 14, and the spacing between the wirings 14. Thus, the number of photodiodes 3 arranged in the slice direction (the number of slice columns) is determined. In the present embodiment, since the wirings 14 are drawn out to the left and right, the number of photodiodes 3 arranged in the slice direction is given by twice the number of wirings 14 that can be formed in the gap.

  Each switching chip 4 includes a plurality of CMOS transistors 5 formed as switching elements on a silicon substrate. The plurality of transistors 5 are connected to the plurality of photodiodes 3, respectively. N / 2 photodiodes 3 arranged on the same slice column on the same side are commonly connected to the signal readout line 7 via N / 2 wirings 14 and N / 2 transistors 5. . A DAS chip 8 is connected to the signal readout line 7.

  When the transistor 5 is turned on, the electric charge accumulated in the photodiode 3 is read out to the signal readout line 7 as a current signal. On / off (gate voltage) of the transistor 5 is controlled by the read control circuit 6. As described above, since N / 2 wirings 14 arranged in the slice direction are commonly connected to the signal readout line 7, by turning on the corresponding N / 2 transistors 5 individually in order, The signals of N / 2 photodiodes 3 can be read out in order. In addition, since the DAS chip 8 is individually connected to each signal readout line 7, the signal readout of all the photodiodes 3 in the module can be performed in the time required to read out N / 2 photodiodes 3 serially. Can be completed.

  When a plurality of transistors 5 corresponding to a plurality of adjacent photodiodes 3 are simultaneously turned on, signals from the plurality of photodiodes 3 can be added in an analog manner. In this case, the plurality of photodiodes 3 constitute one channel. With such a wiring structure and electronic readout control, the slice thickness can be arbitrarily changed.

The plurality of Al wirings 14 connected to the plurality of photodiodes 3 are connected to the plurality of contacts of the rigid printed wiring board 10 via the plurality of metal wires 11 by wire bonding technique. A plurality of bumps 12 formed on the surface of the switching chip 4, the flip chip technique, is connected by soldering to a plurality of bumps corresponding formed on the surface of the rigid printed wiring board 10. Similarly, the plurality of bumps 13 formed on the surface of the DAS chip 8 are connected by soldering to the corresponding plurality of bumps formed on the surface of the rigid printed wiring board 10 by the plip chip technique.

  As described above, by pulling out the Al wiring 14 to the left and right, the number of photodiodes 3 that can be formed in the slicing direction can be doubled as compared with the case where the Al wiring 14 is pulled out only in one direction. Further, by pulling out the Al wiring 14 to the left and right, the wire bonding area between the rigid printed wiring board 10 and the photodiode chip 2 can be expanded, so that the increase in the number of photodiodes 3 can be accommodated. it can.

  Further, by mounting the photodiode chip 2, the switching chip 4, and the DAS chip 8 on one common rigid printed wiring board 10, the connection between these three members is facilitated, and the number of photodiodes 3 is increased. Can handle the increase.

  In addition, since the N / 2 wirings 14 are commonly connected to the signal readout line 7, the number of signal readout lines between the switching chip 4 and the DAS chip 8 can be reduced. Further, the slice thickness can be easily changed by electronic control of the transistor 5 with a simple configuration in which N / 2 wirings 14 are commonly connected to the signal readout line 7.

(Second Embodiment)
In the first embodiment, a plurality of photodiodes are electrically drawn out by a plurality of Al wirings formed on the silicon substrate surface. Therefore, the number of Al wirings, that is, the number (number of columns) of photodiodes arranged in the slice direction is restricted by the gap width of the photodiodes forming the Al wiring.

  In the second embodiment, the above restrictions are alleviated by electrically pulling out a plurality of photodiodes to the back surface by through wirings penetrating the silicon substrate from the front surface to the back surface.

  FIG. 3 is a cross-sectional view of a detector module constituting the radiation detector according to the second embodiment, and FIG. 4 is a cross-sectional view showing in detail the through wiring structure of FIG. On the ceramic rigid printed wiring board 22, at least one photodiode chip 20, at least one switching chip 4, and at least one DAS chip 8 are mounted in common.

  The plurality of bumps 12 formed on the surface of the switching chip 4 are connected to the plurality of bumps formed on the surface of the rigid printed wiring board 20 by soldering by a plip chip technique. Similarly, the plurality of bumps 13 formed on the surface of the DAS chip 8 are connected to the plurality of bumps formed on the surface of the rigid printed wiring board 10 by soldering by the plip chip technique.

  The plurality of photodiodes 5 are formed in a matrix on the surface of the silicon substrate 24. The plurality of photodiodes 5 are respectively connected to the plurality of Al wirings 22 formed on the surface of the silicon substrate 24. The plurality of Al wirings 22 are respectively connected to the plurality of through wirings 23 penetrating from the front surface to the back surface of the silicon substrate 24. The side surfaces of the plurality of through wirings 23 are insulated by, for example, silicon oxide. A plurality of bumps 21 are formed on the rear surface of the silicon substrate 24 at the rear ends of the plurality of through wirings 23. Thus, the plurality of photodiodes 5 are electrically drawn out to the back surface via the plurality of through wirings 23.

  The plurality of bumps 21 formed on the back surface of the silicon substrate 24 are connected to the corresponding plurality of bumps formed on the surface of the rigid printed wiring board 22 by soldering.

  Thus, since the plurality of photodiodes 5 are drawn out to the back surface of the substrate by the plurality of through wires 23, the number (number of columns) of the photodiodes 5 arranged in the slicing direction is released from the above restriction and increased. It becomes possible.

(Third embodiment)
The radiation detector is housed in a gantry housing of the X-ray CT apparatus. The width of the detector module in the slice direction is limited by the size of the internal space of the gantry housing. As described above, when a photodiode chip, a switching chip, and a DAS chip are mounted in common on the surface of a rigid printed wiring board, the mounting area of the photodiode chip is pressed against the switching chip and the DAS chip. Is done. Therefore, the number (number of columns) of photodiodes 5 arranged in the slice direction is restricted by the mounting area of the photodiode chip.

The third embodiment reduces this restriction.
FIG. 5 shows a cross-sectional view of a detector module constituting the radiation detector according to the third embodiment. At least one photodiode chip 20 and at least one switching chip 4 are mounted on the surface of the rigid multilayer wiring board 30. At least one DAS chip 8 is mounted on the back surface of the rigid multilayer wiring board 30. By disposing the DAS chip 8 on the back surface, the width of the detector module in the slice direction can be expanded, and the number of photodiodes 5 (number of columns) arranged in the slice direction can be increased.

  The plurality of photodiodes of the photodiode chip 20 are connected to a plurality of terminals of a printed wiring formed on the surface of the multilayer wiring board 30 via a plurality of through wirings penetrating the silicon substrate from the front surface to the back surface. It is connected to a plurality of transistors of the switching chip 4 through printed wiring.

  The plurality of bumps 12 formed on the surface of the switching chip 4 are connected to the corresponding plurality of bumps formed on the surface of the rigid multilayer wiring board 30 by solder 35 by the plip chip technique.

  The plurality of bumps 13 formed on the surface of the DAS chip 8 are connected to the corresponding plurality of bumps formed on the back surface of the rigid multilayer wiring board 30 by soldering by a plip chip technique.

  FIG. 6 shows a cross-sectional structure of the multilayer wiring board 30. The multilayer wiring board 30 includes a plurality of thin wiring substrates 31 to 34 that are stacked. The plurality of wiring boards 31 to 34 are connected to each other by a plurality of via holes.

  The wiring of each wiring board 31 to 34 and the wiring connection between the plurality of wiring boards 31 to 34 connect the plurality of bumps 12 of the switching chip 4 and the corresponding bump 13 of the DAS chip 8. Designed for.

  Thus, the multilayer wiring board 30 makes it possible to arrange the DAS chip 8 on the back surface. As a result, the mounting area of the photodiode chip pressed by the DAS chip 8 can be expanded, and the number (number of columns) of photodiodes 5 arranged in the slice direction can be increased.

(Fourth embodiment)
FIG. 7 is a cross-sectional view of a detector module according to the fourth embodiment.
In the third embodiment, a photodiode chip and a switching chip are mounted on the surface of the multilayer wiring board, and a DAS chip is mounted on the back surface of the multilayer wiring board, thereby expanding the mounting area of the photodiode chip.

  On the other hand, in the fourth embodiment, only the photodiode chip 20 is mounted on the front surface of the multilayer wiring board 40, and the switching chip 4 is mounted on the back surface of the multilayer wiring board 40 together with the DAS chip 8. The mounting area of the diode chip 20 is further expanded.

  The plurality of photodiodes of the photodiode chip 20 are connected to a plurality of terminals formed on the surface of the multilayer wiring board 40 via a plurality of through wirings penetrating the silicon substrate from the front surface to the back surface.

  The plurality of bumps 12 formed on the surface of the switching chip 4 are connected to the corresponding plurality of bumps formed on the back surface of the rigid multilayer wiring board 40 by soldering by the plip chip technique.

  The plurality of bumps 13 formed on the surface of the DAS chip 8 are connected to the corresponding plurality of bumps formed on the back surface of the rigid multilayer wiring board 40 by soldering by the plip chip technique.

  The wiring of each of the plurality of wiring boards constituting the multilayer wiring board 40 and the wiring connection between the plurality of wiring boards are performed by connecting the plurality of photodiodes of the photodiode chip 20 to the plurality of corresponding switching chips 4. Designed to connect to transistors.

  Thus, by mounting only the photodiode chip 20 on the front surface of the multilayer wiring board 40 and mounting the switching chip 4 and the DAS chip 8 on the back surface of the multilayer wiring board 40, the mounting area of the photodiode chip 20 is reduced. Further enlargement can facilitate the increase in the number (number of columns) of photodiodes 5 arranged in the slice direction.

(Fifth embodiment)
FIG. 8 shows a cross-sectional view of a detector module constituting the radiation detector according to the fifth embodiment. The switching chip 50 and the DAS chip 8 are mounted on a ceramic rigid printed wiring board 52, and the photodiode chip 20 is mounted on the surface of the switching chip 50.

  A plurality of bumps are formed on the front surface of the switching chip 50 and connected to the plurality of bumps on the back surface of the photodiode chip 20 by soldering. A plurality of bumps 51 are also formed on the back surface of the switching chip 50, and are connected to a plurality of transistors on the surface via through wirings. The plurality of bumps 51 on the back surface of the switching chip 50 are connected to the corresponding plurality of bumps formed on the surface of the rigid printed wiring board 52 by solder. Similarly, the plurality of bumps 13 formed on the surface of the DAS chip 8 are connected to the plurality of bumps formed on the surface of the rigid printed wiring board 52 by soldering by the plip chip technique.

  Thus, even if the photodiode chip 20 is mounted on the surface of the switching chip 50, the same effect as that of the third embodiment can be obtained.

(Sixth embodiment)
FIG. 9 shows a cross-sectional view of a detector module constituting the radiation detector according to the sixth embodiment. The switching chip 50 is mounted on the surface of the rigid multilayer wiring board 60, the DAS chip 8 is mounted on the back surface of the rigid multilayer wiring board 60, and the photodiode chip 20 is mounted on the surface of the switching chip 50. Yes.

  A plurality of bumps are formed on the front surface of the switching chip 50 and connected to the plurality of bumps on the back surface of the photodiode chip 20 by soldering. A plurality of bumps 51 are also formed on the back surface of the switching chip 50, and are connected to a plurality of transistors on the surface via through wirings. The plurality of bumps 51 on the back surface of the switching chip 50 are connected to the corresponding plurality of bumps formed on the surface of the multilayer wiring board 60 by solder. The plurality of bumps 13 formed on the surface of the DAS chip 8 are connected to the plurality of bumps formed on the back surface of the multilayer wiring board 60 by soldering using a plip chip technique.

  The wiring of each of the plurality of wiring boards constituting the multilayer wiring board 60 and the connection of the wiring between the plurality of wiring boards are performed by connecting the plurality of transistors of the switching chip 4 to the plurality of bumps of the corresponding DAS chip 8 respectively. Designed to connect.

  Thus, even when the photodiode chip 20 is mounted on the front surface of the switching chip 50 and the DAS chip 8 is mounted on the back surface of the multilayer wiring board 60, the same effects as in the fourth embodiment can be obtained. .

  As described above, the first to sixth embodiments can be developed to various applications described below, and can realize specific and significant configurations.

(Seventh embodiment)
FIG. 10 shows a configuration of an X-ray CT apparatus according to the seventh embodiment. The rotating ring 102 is driven by the gantry driving unit 107 at a high speed rotation of 1 second or less per rotation. An X-ray tube 101 for generating cone beam (quadrangular pyramid) or fan beam-like X-rays with respect to the subject P placed in the effective visual field region FOV is attached to the rotating ring 102. ing. The X-ray tube 101 is supplied with power necessary for X-ray exposure from the high voltage generator 109 via the slip ring 108. Thereby, the X-ray tube 101 is a so-called cone beam X-ray that spreads in two directions, ie, a channel direction C orthogonal to the body axis direction of the subject and a slice direction A (= direction parallel to the rotation axis) orthogonal thereto. Fan beam X-rays are generated.

  Further, a radiation detector 103 for detecting X-rays transmitted through the subject P is attached to the rotating ring 102 in a direction facing the X-ray tube 101. The radiation detector 103 includes a plurality of (for example, 38) detector modules. FIG. 11 shows a development view of one detector module. The detector module 1030 includes a scintillator and a photodiode chip having a plurality of detection elements including photodiodes 1031 and 1032. The plurality of detection elements 1031 and 1032 are arranged in a matrix in the two directions of the channel direction C and the slice direction S. In the X-ray CT apparatus according to the present embodiment, the plurality of detector modules 1030 are not planar but are arranged along an arc centered on the focal point of the X-ray tube 101.

  The detector module 1030 includes a switching chip and a DAS chip together with the photodiode chip having the plurality of detection elements 1031 and 1032 as described above. These photodiode chip, switching chip, and DAS chip are mounted on a single rigid printed wiring board.

  One detection element 1031 has a sensitive area with a width of 0.5 mm in the slice direction and a width of 1 mm in the channel direction. The other detection element 1032 has a sensitive area with a width of 1 mm in the slice direction and a width of 1 mm in the channel direction.

  For example, 16 detection elements 1031 having a width of 0.5 mm are arranged in the slice direction C. The 16 detection elements 1031 arranged in the slice direction C are referred to as a first detection element array group 1033. In addition, with respect to the slice direction S, the detection elements 1032 having a width of 1 mm are arranged on each side of the first detection element array 1033 in a plurality, for example, twelve, which is smaller than the number of detection elements 1031 arranged. The twelve detection elements 1032 arranged in the slice direction C are referred to as a second detection element array group 1034.

  In the present embodiment, the number (for example, 16) of the detection elements 1031 arranged in the slice direction C is larger than the number (for example, 12) of the detection elements 1032 arranged on both sides thereof, and the total number (for example, It is designed to be less than 24).

  M × N detected by such a radiation detector 103 (M = 24 rows × 38 = 912 in the above example, N = 40 (= 16 columns + 2 × 12 columns)) A huge amount of data relating to all channels (data for M × N channels per behaviour is hereinafter referred to as “two-dimensional projection data”) is once collected in a chip data acquisition circuit (DAS) 104, and The data is transmitted to a data processing unit to be described later via a contactless data transmission device 105 that applies optical communication.

  It should be noted that the detection operation by the radiation detector 103 is repeated, for example, about 1000 times during one rotation (about 1 second), so that an enormous amount of 2D projection data for M × N channels is obtained for 1 second (1 rotation). The data collection circuit 104 and the non-contact data transmission device 105 are designed to be ultra-high-speed processing in order to transmit such enormous and high-speed generated 2D projection data without time delay. .

  The data processing unit is centered on the host controller 110, and includes a preprocessing device 106 that performs preprocessing such as data correction, a storage device 111, an auxiliary storage device 112, a data processing device 113, a reconstruction device 114, an input device 115, and a display device. 116 are interconnected via a data / control bus 300. Further, an external image processing device 200 including an auxiliary storage device 201, a data processing device 202, a reconstruction device 203, an input device 204, and a display device 205 is connected via the bus 300.

  Below, the effect about the X-ray CT apparatus used as the said structural example is demonstrated. In the following, volume data composed of a plurality of voxel data is reconstructed based on the subject transmission X-ray data detected by the radiation detector 103, and from this, a tomographic image of an arbitrary cross section is generated from an arbitrary direction. A case (operation) of reconstructing various images such as a projected image and a three-dimensional surface display image will be described.

  FIG. 12 shows data processing and its flow in the X-ray CT apparatus of this embodiment. The X-ray transmitted through the subject is converted into two-dimensional projection data of an analog electrical signal by the radiation detector 103, and further converted into two-dimensional projection data of a digital electrical signal by the data acquisition circuit 104, and then contactless data transmission. Via the device 105, it is sent to a preprocessing device 106 that performs various corrections, and receives sensitivity correction and the like.

  In the following, the width of the sensitive area of the photodiode is defined as a converted value on the rotation center axis of the X-ray tube. That is, “a photodiode having a sensitive area width of 1 mm” means “a photodiode having a sensitive area width corresponding to 1 mm on the rotation center axis of the X-ray tube”, and X-rays diffuse radially. The width of the actual sensitive area of the photodiode is less than 1 mm according to the ratio of the actual distance between the X-ray focal point and the photosensitive area of the photodiode with respect to the distance between the X-ray focal point and the rotation center axis. Slightly wider.

  Here, the conversion action to the analog electric signal in the radiation detector 103 is the first detection element row 1033 including the detection elements 1031a having a width of 0.5 mm for the X-rays detected in the vicinity of the center in the slice direction A. The remaining portion is performed by the second detection element array 1034 configured by the detection element 1032 having a larger width of 1 mm. In other words, according to the first detection element array 1033, the two-dimensional projection data collection and the conversion to an analog electric signal are achieved while maintaining a higher resolution than the remaining portion. .

  Further, as described above, the X-ray tube 101 in the X-ray CT apparatus of the present embodiment generates cone beam-shaped or fan beam-shaped X-rays. In such an X-ray tube 101, a collimator having an appropriate configuration is installed in the vicinity of the X-ray tube 101, and the like. It is possible to change the X-ray beam thickness, that is, the “slice width (= size obtained by bundling a plurality of uniform“ slice thicknesses ”)” with respect to the subject.

  In FIG. 13, the detection elements used when photographing 16 slices having a thickness of 0.5 mm are indicated by hatching. In FIG. 14, the detection elements used when 16 slices having a thickness of 1 mm are photographed are indicated by hatching. Further, in FIG. 15, detection elements used when 32 slices having a thickness of 1 mm are photographed are indicated by hatching.

  When 16 slices having a thickness of 0.5 mm are photographed, data capable of reconstructing 16 tomographic images having a thickness of 0.5 mm are collected by individually reading electrical signals from the detection element 1031 having a width of 0.5 mm. be able to.

  In addition, when 16 slices having a thickness of 1 mm are photographed, an electric signal of the detection element 1031 having a width of 0.5 mm between adjacent pairs is read out simultaneously, so that the detection element 1031 of the adjacent pair having a width of 0.5 mm is a single element. Data that can reconstruct 16 tomographic images of 1 mm thickness by individually reading out the electrical signals of a total of 8 1 mm-wide detection elements 1032, 4 on each side near the center. Can be collected.

  When 32 slices having a thickness of 1 mm are photographed, the electric signals of the detection elements 1031 having a width of 0.5 mm between adjacent pairs are read out simultaneously, so that the detection elements 1031 of the adjacent pairs having a width of 0.5 mm can be obtained as a single element. In addition, data that can reconstruct 32 tomograms having a thickness of 1 mm can be collected by individually reading the electrical signals of the detection element 1032 having a width of 1 mm.

  Of course, by changing the signal reading method, the slice thickness can be changed variously, such as 2 mm and 3 mm. The number of slices can be arbitrarily changed from 1 to 40. Furthermore, it is also possible to collect data of a plurality of types of slice thicknesses simultaneously. Furthermore, it is possible to collect 8 slices at any slice thickness of 0.5 mm, 1 mm, 2 mm, 3 mm, and 4 mm.

  As described in the first embodiment, such signal read control is performed by connecting a plurality of photodiodes on the same slice column to a single common signal line via a plurality of transistors. / Off control can be realized.

  In the above description, the first detection element array 1033 is used at the time of 0.5 mm slice, but instead of this, “8” of the first detection element array 1033 at the 0.5 mm slice is used. It may be configured to use only the “column”. In addition, various cases can be considered at the time of each slice. In any case, a mode using “8 columns” other than 16 columns or “4 columns” in some cases may be used.

  Now, the 360 ° two-dimensional projection data subjected to sensitivity correction, X-ray intensity correction, and the like in the preprocessing device 106, that is, 1000 sets of two-dimensional projection data, are temporarily stored in the storage device 111 directly. Thereafter, the two-dimensional projection data is sent to the reconstruction device 114, where reconstruction based on the data is performed by a reconstruction algorithm called the Feldkamp method.

  The Feldkamp reconstruction method treats a wide target region in the slice direction A as an aggregate of a plurality of voxels, and uses three-dimensional distribution data (hereinafter referred to as “volume data (a plurality of voxel data is three-dimensional (three-dimensional) This is an approximate reconstruction method that is improved based on the fan beam convolution back projection method. That is, in the Feldkamp reconstruction method, the data is considered as fan projection data and convolved, and the back projection is performed along an oblique ray corresponding to the actual cone angle with respect to the rotation center axis. Note that the fan beam convolution back projection method that has not been improved assumes that the ray is orthogonal to the rotation center axis in back projection, and thus artifacts appear strongly.

  Therefore, by adopting the Feldkamp reconstruction method, a detector wide in the slice direction can be used effectively.

  The reconstructed volume data is directly or once stored in the storage device 111 and then sent to the data processing device 113. Based on the instructions of the operator, the tomographic image of an arbitrary cross section, which is already widely used It is converted into so-called pseudo three-dimensional image data such as a projected image from a direction and a three-dimensional surface image of a specific organ by rendering processing, and is displayed on the display device 116.

  The operator can select and set an arbitrary display form from the tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, a three-dimensional surface image, and the like according to the purpose of inspection / diagnosis. In this case, an image in a different form is generated from one volume data and displayed. In addition, when displaying, not only one type of image but also a mode in which a plurality of types of images are displayed at the same time can be switched to a mode in which one image is displayed according to the purpose. Yes.

  Incidentally, the above-mentioned “tomographic image of arbitrary cross section” means not only a cross section (axial cross section) AX orthogonal to the body axis obtained by a conventional X-ray CT apparatus, but also a sagittal cross section SA, as shown in FIG. It refers to a tomographic image of a cross section orthogonal to the axial cross section AX, such as a coronal cross section CO, and an oblique cross section OB inclined with respect to these cross sections AX, SA, and CO. These extract voxel data of the designated thickness from the volume data, and display them in a bundle.

  The “projected image from an arbitrary direction” refers to, for example, picking up the maximum value of voxel data arranged in the direction set as the arbitrary direction with respect to the volume data, and integrating the arranged voxel data. By taking a value or the like, this is displayed as a two-dimensional image. Furthermore, the “three-dimensional surface image” is a method of, for example, extracting a surface based on a set threshold value and displaying the surface three-dimensionally by shading with a set light source. In this case, the internal structure can be grasped by observing while changing the threshold value.

  As described above, in the X-ray CT apparatus according to the present embodiment, various images can be acquired through a single volume data. By adopting the radiation detector 103, the following is based on the above description. It is possible to enjoy the effects described in 1. That is, the size of the voxel data constituting the volume data varies depending on the system geometry, the data collection speed, and the like, and also greatly depends on the size of the detection elements 1031 and 1032 constituting the radiation detector 103. In this regard, according to the radiation detector 103 in the present embodiment, the detection elements 1031 having small widths are provided in multiple rows (16 rows in the above example), so that for each voxel data, for example, a minimum of 0.5 mm. It is possible to carry out a relatively wide range of imaging while achieving about × 0.5 mm × 0.5 mm. That is, a wide range and high resolution are maintained.

  On the other hand, since the radiation detector 103 of the present embodiment has a multi-row structure of the detection elements 1031 and 1032, large volume data can be obtained by one rotation. Isotropic voxel data can be collected. Therefore, the resolution of the tomographic image of an arbitrary cross section can be made substantially equal, and a clinically useful image can be acquired. In other words, since data can be collected with isotropic voxels for all of the X-axis, Y-axis, and Z-axis, the tomographic images of the axial section AX, sagittal section SA, and coronal section CO are expressed with the same resolution. Diagnosis based on the obtained image can be performed. Clinically, so-called isotropic can be realized based on a slice of 0.5 mm head and 1 mm abdomen.

  The fact that “wide range” imaging is possible implies that “rapid” imaging is possible, that is, it is possible to reduce the total amount of exposure to the subject. However, apart from that, according to the radiation detector 103 in the present embodiment, the number (or the number of columns) of the detection elements 31a having a small width is made smaller than that of the detection elements 31 having a large width. Therefore, for example, the number of data that must be handled is enormous as in the case of miniaturization of the detection element in the two-dimensional array type detector disclosed in JP-T-8-509896 as described in the prior art. From this point, it can be pointed out that “rapid” processing is possible. Of course, it is obvious that the “crosstalk” problem pointed out with respect to the publication is not a big problem in this embodiment.

  In addition, according to the radiation detector 103 in the present embodiment, there is a specific effect for imaging in which the X-ray tube 101 and the radiation detector 103 are continuously rotated.

  First, according to the scan of one rotation, by performing the above data processing, as described above, from the two-dimensional projection data from multiple directions obtained by only one rotation, an object region wide in the slice direction A is obtained. One volume data having no time difference in the slice direction A can be obtained. In addition to the axial cross section AX, it is possible to observe a tomographic image at a certain time (same time). As described above, the form of the display image in this case can be selected and set from a tomographic image having an arbitrary cross section, a projection image in an arbitrary direction, a three-dimensional surface image, and the like.

  According to continuous rotation scanning, when the same processing as in the case of one rotation is repeatedly performed on two-dimensional projection data from multiple directions obtained by a plurality of rotations, one volume data is obtained. Rather than multiple. Even when reconfiguration is performed for each rotation, the same number of sets as the number of rotations can be obtained, and the range of data used for reconstruction (range of the rotation angle of the system) is gradually shifted, so that More volume data that is slightly different can be obtained.

  As for the form of the display image in the case of this continuous rotation scan, as in the case of one rotation, depending on the setting of the operator, a tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, a three-dimensional surface image, etc. It is possible to select from.

  Further, by generating slightly different images in time in the set display form from the volume data slightly different in time, and displaying the images in order, the operator can display the images. As shown in FIG. 17, it becomes possible to observe a moving image in real time. That is, it is possible to display an image as a moving image in parallel with continuous scanning.

  In the following, data processing in the case of this continuous rotation scan will be described a little. However, in the following description, the angle range of projection data necessary to reconstruct one three-dimensional image data will be described as 360 °. As already described, the angle range is an angle other than 360 °, For example, it may be 180 ° + view angle.

  First, the X-ray tube 101 continuously rotates at a high speed around the subject together with the radiation detector 103. The time required for one rotation is t0 (1 second in the above example). The projection data collected one after another is preprocessed in almost real time. Then, the reconstruction device 114 reconstructs the volume data “I” based on the pre-processed 360 ° projection data. Then, based on the reconstructed volume data I, the data processing device 113 generates image data “DI” such as a tomographic image of an arbitrary cross section, a projection image from an arbitrary direction, and a three-dimensional surface image. The image data “DI” is displayed on the display device 116.

  When displaying a moving image as described above, for each scan from the first scan to the n-th scan, a series of processing from the scan to the image display is performed in parallel (for example, the first scan In parallel with the reconstruction processing based on the image, the second scan is performed, etc.), and the images are reconstructed one after another based on the two-dimensional projection data obtained in succession, and displayed one after another. .

  For this reason, the reconstruction device 114 has a volume shorter than the time t0 required to collect projection data for a predetermined angle range (here, 360 °) in parallel with the two-dimensional projection data collection operation (scan). The processing capability necessary to reconstruct the data I is provided. Further, the data processing device 113 has a processing capability necessary for generating the image data DI from the volume data I in a shorter time than the reconstruction time of the volume data I. Further, the display device 116 is equipped with a counter, a memory, and the like necessary for starting display of the image data DI from a starting point Ts or an ending point Te during the period of the operation of collecting projection data originating from the image data DI after a predetermined time. is doing.

  By the way, regarding the point just described, it is naturally required that the signal processing capability around the radiation detector 103 is also high. In this regard, according to the radiation detector 103 in the present embodiment, as described above, since “rapid” signal processing is possible, it greatly contributes to smooth execution of the moving image display as described above. Moreover, it can be achieved while maintaining a high resolution over a wide range.

  In the above embodiment, data processing and display operations such as reconstruction and cross-section conversion are performed in the X-ray CT apparatus 100 (such a form is common), but in the present invention. Instead, these data processing and the like may be executed in the external image processing apparatus 200 in FIG. When such an external image processing apparatus 200 is used, the data sent from the X-ray CT apparatus 100 to the image processing apparatus 200 may be before reconstruction, after reconstruction, or immediately before display after data processing. In any state, the effect of the above-described embodiment is not disturbed.

  Further, in the above description, it is possible to display a tomographic image of an arbitrary cross section, a projection image in an arbitrary direction, a three-dimensional surface image, and the like. It is good also as a structure which displays the time change information, such as a value and an electrocardiogram, with a graph, and also displays the time of the main image currently displayed on the graph.

  Furthermore, this embodiment can be implemented with various modifications. For example, the above-described embodiment may be applied to a helical scan in which at least one of the gantry and the bed is moved during scanning so that the X-ray tube 101 draws a spiral trajectory around the subject. At this time, if the imaging condition including at least one of the detection element array, helical pitch, scan range, scan time, and tube current used for data collection is set to an optimum value, volume data consisting of isotropic voxel data is obtained. be able to. In realizing this isotropic, it is desirable from the viewpoint of high resolution to collect data using only the first detection element array having a narrow detection width. For example, the imaging conditions are set such that the imaging area head 18 cmφ, 0.5 mm detection element array × 4 arrays, helical pitch 3, scan range 60 mm, scan time 20 seconds, 150 mAs, and reconstruction pitch 0.3 mm.

  In addition, the arrangement | sequence of a detection element and the structure of a detector are either or any combination of 1st-6th embodiment.

(Eighth embodiment)
FIG. 18 shows a schematic configuration of an X-ray CT apparatus according to the eighth embodiment. In FIG. 18, an X-ray CT apparatus 2010 includes a gantry 2011 that collects projection data of a subject, and a data processing system unit that performs image reconstruction processing, reconstruction image display, and the like based on the collected projection data. It is configured.

  The gantry 2011 includes an X-ray tube 2014, a slit 2016, a bed for placing a subject 2018, a diagnostic opening for inserting a subject to make a diagnosis, a gantry driving unit 2038, and a radiation detection system 2020. ing.

  The X-ray tube 2014 is a vacuum tube that generates X-rays. The X-ray tube 2014 generates X-rays by accelerating electrons by a high voltage generated by a high-voltage generator 2012 described later and colliding with a target.

  The slit 2016 is provided between the X-ray tube 2014 in the gantry 2012 and the subject, shapes the cone-shaped X-ray beam exposed from the X-ray focal point of the X-ray tube 2014, and has a required solid angle. The X-ray beam is formed.

  The bed 2018 can be sliced along the body axis direction of the subject by driving the bed driving unit.

  The gantry drive unit 2038 performs drive control such as rotating the X-ray tube 2014 and the radiation detection system 2020 integrally around a central axis parallel to the body axis direction of the subject inserted into the diagnostic aperture. Do.

  The radiation detection system 2020 is a system including a radiation detector 2025 and a data collection device (hereinafter referred to as DAS) 2030. The radiation detector 2025 includes a plurality of detector modules arranged. The radiation detection system 2020 according to the present embodiment is sufficient if the conventional detector channel direction resolution is about 0.5 to 1 mm, so that the slice axis direction also has a resolution equivalent to the channel direction. A 0.5 mm slice is described as the minimum imaging slice thickness. The DAS 2030 is composed of a plurality of DAS chips, and performs control for bundling the X-ray transmission data sent from the radiation detector 2025 into an amplification process, an A / D conversion process, and a predetermined slice thickness. Send to processing unit 2032. The DAS chip is sometimes called DAM-ASSY.

  Furthermore, the radiation detection system 2020 according to the present embodiment has the characteristics described later in order to significantly reduce noise.

  The data processing system unit includes a high voltage generation device 2012, a data processing device 2032, a storage device 2034, a host controller 2036, an input device 2040, a reconstruction device 2042, and a display device 2044.

  The high voltage generator 2012 is a device that supplies a high voltage to the X-ray tube 2014, and includes a high voltage transformer, a filament heating converter, a rectifier, a high voltage switch, and the like. High voltage supply to the X-ray tube 2014 by the high voltage generator 2012 is performed by, for example, a contact-type slip ring mechanism.

  The host controller 2036 is equipped with a computer circuit having a CPU, is connected to the high voltage generator 2012, and has a bed driving unit, a gantry driving unit 2038, a radiation detection system 2020 (not shown) in the gantry 2011 via a bus. Are connected to each. In addition, the host controller 2036, the data processing device 2032, the storage device 2034, the reconstruction device 2042, the display device 2044, and the input device 2040 are interconnected via a bus, and image data and control data are mutually connected at high speed via the bus. Etc. can be delivered.

  For example, the host controller 2036 performs control as described below to perform X-ray transmission data (projection data) collection processing. That is, the host controller 2036 stores the scan conditions such as slice thickness input from the operator via the input device 2040 in the internal memory, and the stored scan conditions (or scans directly set by the operator in the manual mode). Of the high voltage generator 2012, the bed driving unit, the gantry driving unit 2038, and the bed 2018 in the body axis direction, the feeding speed, the gantry 2011 (the X-ray tube 2014 and the radiation detection system 2020). The high voltage generator 2012, the bed driving unit, and the gantry driving unit 2038 are driven while controlling the rotation speed, the rotation pitch, the X-ray exposure timing, and the like. Then, a cone-shaped X-ray beam is irradiated from a plurality of directions to a desired imaging region of the subject, and transmitted X-rays transmitted through the imaging region of the subject are transmitted through each detection element of the radiation detection system 2020 as X. It can be detected as line transmission data.

  At the same time, the host controller 2036 controls on / off of the switching element of the radiation detection system 2020 based on the scan condition (or the manual mode scan condition) stored in the internal memory. The host controller 2036 switches the connection state between each detection element (photodiode) included in the radiation detection system 2020 and the DAS, and bundles X-ray transmission data detected by each detection element in a predetermined unit. Then, it is sent to the DAS as X-ray transmission data of a plurality of slices corresponding to the scan condition, and a predetermined process is executed.

  The data processing device 2032 includes a computer circuit having, for example, a CPU, and holds projection data for, for example, 16 slices collected by each DAS chip of the radiation detection system 2020. The data processing device 2032 adds the projection data of the same slice obtained from the multi-directions by the rotation of the gantry 2011 described above or the multi-directional data obtained by the addition processing as necessary. Interpolation processing, correction processing, etc. are performed.

  The storage device 2034 stores data necessary for data processing in the data processing device 2032.

  The reconstruction device 2042 reconstructs projection data obtained by data processing by the data processing device 2032 according to the Feldkamp reconstruction method, and generates reconstructed image data for 16 slices.

  The display device 2044 displays the reconstructed image data generated by the reconstruction device 2036.

  The input device 2040 includes a keyboard, various switches, a mouse, and the like, and can input various scan conditions such as slice thickness and the number of slices through an operator.

  The reconstruction device 2042 has a large-capacity auxiliary storage device capable of storing the generated reconstructed image data.

  Next, the radiation detection system 2020 according to the present embodiment will be described in detail.

  Each detector module of the radiation detection system 2020 includes one detector block and one DAS block. That is, the radiation detector 1020 is configured by configuring a detector module with one detector block and one DAS block and arranging the detector modules in the channel direction. Hereinafter, each component will be described in order.

(Detector module and radiation detector)
FIG. 20A shows a top view of the detector block 2200. FIG. The detector block 2200 includes a photodiode chip 2201 that outputs light converted from X-rays by a scintillator into electrical signals by a plurality of photodiodes, and a DAS chip by bundling collected X-ray transmission data in predetermined units. A switching chip 2202 made up of a group of CMOS (Complementary MOS) is mounted on a rigid multilayer wiring board 2220. A connector electrically connected to the switching chip 2202 is attached to the back surface of the multilayer wiring board 2220.

  FIG. 19 shows a development view of the photodiode chip 2201. As shown in the figure, the photodiode chip 2201 has a plurality of photodiodes 2001 and 2002. One photodiode 2002 has a substantially square sensitive area of 1 mm in the channel direction and 1 mm in the slice direction. The other photodiode 2001 has a sensitive area of 0.5 mm, which is the same 1 mm as that of the photodiode 2002 in the channel direction, and ½ that of the photodiode 2002 in the slice direction.

  Sixteen photodiodes 2001 having a width of 0.5 mm are arranged in the slice direction. For the channel direction, 48 are arranged. Twelve 1 mm wide photodiodes 2002 are arranged on each side of the 0.5 mm wide photodiode 2001 in the slice direction. Similarly, 48 channels are arranged in the channel direction.

(DAS block and data collection device)
FIG. 20B shows a top view of the DAS block 2300. The DAS block 2300 has a rigid printed wiring board 2301. On the front surface of the printed wiring board 2301, a connector 2302 that can be attached to and detached from the connector formed on the back surface of the detector block 2200 is formed. A plurality of DAS chips 2303 are mounted on the printed wiring board 2301.

(Detector module and radiation detector)
Next, the detector module 2022 will be described.

  FIG. 20 (c) shows a top view of the detector module 2022. As shown in FIG. 20 (c), the detector module 2022 includes one detector block 2200 and one DAS block 2300. A fixed and electrical connection between the detector block 2200 and the DAS block 2300 is made by a connector 2302. The DAS block 2300 performs processing for collecting X-ray image data detected by the detector block 2200 connected to the connector 2302. Therefore, a single detector module 2022 is a system that performs independent data detection and data collection. The radiation detection system 2020 is configured by arranging a plurality of detector modules 2022 shown in FIG. 20C along the channel direction.

  If one of the connected detector block 2200 and DAS block 2300 fails, they can be separated and replaced with a normal block.

  FIG. 21A shows a cross-sectional view of the detector module 2022. The detector module 2022 relays between the detector block 2200 and the DAS block 2300 with a rigid printed wiring board 2210. The detector block 2200 is mounted on the surface of the relay substrate 2210 via the connector 2214, and the DAS block 2300 is mounted outside the X-ray irradiation area on the back surface via the connector 2302. Further, as shown in FIG. 21B, the DAS block 2300 may be mounted outside the X-ray irradiation area on the surface of the relay substrate 2210.

  In this manner, by disposing the DAS block 2300 outside the X-ray passage region, it is possible to prevent a DAS malfunction due to the ionization effect of X-rays.

FIG. 22 shows a cross-sectional view of the detector module 2022 in which the DAS block 2300 is provided on both surfaces of the relay substrate 2210. Note that the detector module 2022 shown in FIG. 22 has a support 2212 for arranging a plurality of detector blocks 2200 and heat dissipation for dissipating heat from the DAS chip 2303 for alignment and stability between the blocks. It has a fin / pin 2320, a heat conductive rubber 2322 for connecting the radiating fin / pin 2320 and the DAS chip 2303, and a fan 2324 for cooling the data collection device 2030.

FIG. 23A and FIG. 23B are cross-sectional views of the detector module 2022. The detector module 2022 is formed by laminating a detector block 200, a DAS block 2300, and a rigid relay substrate 2210 made of ceramic or glass epoxy resin (glass epoxy resin) . Further, the detector block 2200 is formed by laminating a scintillator 2201, a photodiode chip 2204, and a switching chip 2202. The DAS block 2300 is obtained by sealing a predetermined number of DAS chips 2303 with resin. The scintillator 2201 and the photodiode chip 2204 are sealed with a resin 2206.

  According to this detector module 2022, since the DAS block 2300, the relay board 2210, and the detector block 2200 are stacked, it is possible to reduce the size of the apparatus. In this example, the DAS block 2300 is configured to exist in the X-ray passing area. Therefore, it is more preferable to provide lead on the X-ray incident side of each DAS chip 2303.

  In addition, since the detector module 2022 has the DAS block 2300 arranged outside the X-ray irradiation region, it is possible to prevent malfunction due to the ionization effect of X-rays.

  Of course, the assumed positions of the detector block 2200 and the DAS block 2300 are not limited. For example, as shown in FIG. 24A, the detector block 2200 is arranged at the center of the relay board 2210, and the DAS block 2300 is arranged at the edge of the relay board 2210 as shown in FIG. It may be.

  In this embodiment, the following effects are produced.

  First, because the detector block 2200 and the DAS block 2300 are mounted on a rigid board in common, the flexible PC board is vibrated due to rotation during scanning, disconnection of the connector, the antenna effect of the flexible PC board, etc. Generation of noise can be prevented. Moreover, since a flexible PC board is not used, it is easy to install a shield for noise reduction. As a result, it is possible to achieve a significant noise reduction compared to the conventional case.

  Second, since the detector block and the DAS block are connected by connector connection, the space required for connection between the devices is small, and the device can be downsized. In addition, since the connection and removal of the device are easy by the connector connection, it is possible to easily replace the device and the like compared to the bump connection.

  Third, since the detector module and the DAS block are in a one-to-one correspondence, for example, only a part of the devices can be easily replaced. Therefore, it is not necessary to replace the entire data collection device, and the running cost can be reduced.

(Ninth embodiment)
FIG. 25 shows a configuration of an X-ray CT apparatus according to the ninth embodiment. The X-ray CT apparatus 3010 includes a gantry 3012 that collects projection data of a subject, and a system unit 3014 that performs image reconstruction processing, reconstruction image display, and the like based on the collected projection data.

  The gantry 3012 includes an X-ray tube 3020, a slit 3022, a bed 3024 for placing a subject, a diagnostic opening for inserting a subject to make a diagnosis, a gantry driving unit 3026, and a radiation detector 3028. ing.

  The X-ray tube 3020 is a vacuum tube that generates X-rays. The X-ray tube 3020 generates X-rays by accelerating electrons by a high voltage generated by a high-voltage generator 3030 described later and colliding with a target.

  The slit 3022 is provided between the X-ray tube 3020 in the gantry 3012 and the subject, shapes the cone-shaped X-ray beam exposed from the X-ray focal point of the X-ray tube 3020, and has a required solid angle. The X-ray beam is formed.

  The bed 3024 can be sliced along the body axis direction of the subject by driving a bed driving unit (not shown).

  The gantry drive unit 3026 is a drive that rotates the X-ray tube 3020 and the radiation detector 3028 together around a central axis parallel to the body axis direction of a subject inserted in a diagnostic aperture (not shown). Take control.

  The radiation detector 3028 is formed by arranging a plurality of detector modules in the channel direction.

  The system unit 3014 includes a high voltage generation device 3030, a host controller 3031, a data processing device 3032, a storage device 3034, a reconstruction device 3036, a display device 3038, an input device 3040, and an auxiliary storage device 3042.

  The high voltage generator 3030 is a device that supplies a high voltage to the X-ray tube 3020, and includes a high voltage transformer, a filament heating converter, a rectifier, a high voltage switch, and the like. Application of a high voltage to the X-ray tube 3020 by the high voltage generator 3030 is performed by, for example, a contact-type slip ring mechanism.

  The host controller 3031 is equipped with a computer circuit having a CPU, is connected to the high voltage generator 3030, and has a bed driving unit, a gantry driving unit 3026 (not shown) in the gantry 3012 via the bus B, a radiation detector. 3028, respectively. In addition, the host controller 3031, the data processing device 3032, the storage device 3034, the reconstruction device 3036, the display device 3038, the input device 3040, and the auxiliary storage device 3042 are interconnected via the bus B, and are mutually connected through the bus B. It is configured so that image data, control data, etc. can be transferred at high speed.

  For example, the host controller 3031 executes control as described below to perform X-ray transmission data (projection data) collection processing. That is, the host controller 3031 stores the scan conditions such as slice thickness input from the operator 3000 via the input device 3040 in the internal memory, and the stored scan conditions (or is set directly from the operator 3000 in the manual mode). Based on the scanning conditions), the feeding amount in the body axis direction of the high voltage generator 3030, the bed driving unit, the gantry driving unit 3026, and the bed 3024 (not shown), the feeding speed, the gantry 3012 (the X-ray tube 3020 and the radiation detection) The high-voltage generator 3030, the bed driving unit (not shown), and the gantry driving unit 3026 are driven while controlling the rotation speed, rotation pitch, X-ray exposure timing, and the like. Then, a cone-shaped X-ray beam is irradiated from a plurality of directions onto a desired imaging region of the subject P, and transmitted X-rays transmitted through the imaging region of the subject are passed through each detection element of the radiation detector 3028. It can be detected as X-ray transmission data.

  At the same time, the host controller 3025 performs switching control of each switch group of the radiation detector 3028 based on the scan conditions (or manual mode scan conditions) stored in the internal memory. The host controller 3025 switches the connection state between each detection element (photodiode) included in the radiation detector 3028 and the DAS, and bundles X-ray transmission data detected by each detection element in a predetermined unit. Then, a plurality of slices of X-ray transmission data corresponding to the scanning conditions are sent to the data acquisition element, and predetermined processing is executed.

  The data processing device 3032 includes a computer circuit having, for example, a CPU, and holds projection data for 32 slices collected by each data collection element of the radiation detector 3028. Then, the data processing device 3032 performs processing for adding all the projection data of the same slice obtained from multiple directions by rotation of the gantry 3012 described above, and multidirectional data obtained by the addition processing as necessary. Interpolation processing, correction processing, etc. are performed.

  The storage device 3034 stores data necessary for data processing in the data processing device 3032.

  The reconstruction device 3036 reconstructs projection data obtained by data processing by the data processing device 3032 according to the Feldkamp reconstruction method, and generates reconstructed image data for 8 slices.

  The display device 3038 displays the reconstructed image data generated by the reconstruction device 3036.

  The input device 40 includes a keyboard, various switches, a mouse, and the like, and can input various scanning conditions such as a slice thickness and the number of slices through an operator.

  The auxiliary storage device 3042 is a device having a large-capacity storage area that can store the reconstructed image data generated by the reconstruction device 3036.

  FIG. 26A is a cross-sectional view of the detector module 3280. A photodiode chip 3283 is disposed on the back surface of the scintillator 3281 and is attached with an adhesive 3282. The switching chip 3285 is bump-connected to the back surface of the photodiode chip 3283. The switching chip 3285 is mounted on the surface of a rigid multilayer wiring board 3287. The switching chip 3285 and the rigid multilayer wiring board 3287 are electrically connected by solder bumps. The laminated structure of the photodiode chip 3283, the switching chip 3285, and the multilayer wiring board 3287 is sealed with a resin 3284.

  A DAS chip 3289 is mounted on the back surface of the rigid multilayer wiring board 3287. The multilayer wiring board 3287 and the DAS chip 3289 are electrically connected by a plip chip.

  FIG. 27 is a development view of the photodiode chip 3283. The photodiode chip 3283 has a plurality of photodiodes 3001 and 3002. One photodiode 3002 has a substantially square sensitive area of 1 mm in the channel direction and 1 mm in the slice direction. The other photodiode 3001 has a sensitive area of 0.5 mm, which is the same 1 mm as that of the photodiode 3002 in the channel direction, and ½ that of the photodiode 3002 in the slice direction.

  Thirty-two photodiodes 3001 having a width of 0.5 mm are arranged in the slice direction. For the channel direction, 48 are arranged. Eight photodiodes 3002 having a width of 1 mm are arranged on both sides of the photodiode 3001 having a width of 0.5 mm in the slice direction. Similarly, 48 channels are arranged in the channel direction.

  The scintillator 3281 includes a 0.5 mm wide scintillator element and a 1 mm wide scintillator element arranged in the same pattern as the photodiodes 3001 and 3002. A lead separator is fitted between adjacent scintillator elements to prevent crosstalk.

  Each scintillator element is box-shaped (hexahedral), and a light reflecting agent (not shown) is provided in layers on the X-ray incident surface and the slice thickness direction end surface. A photodiode is bonded to the scintillator element on the side of the fluorescence output surface (the surface facing the X-ray incident surface) so as to receive fluorescence through a bonding member such as an adhesive 3282, for example.

  Returning to FIG. The photodiode chip 3283 is composed of the same number of photodiodes as the scintillator elements. Each photodiode and each scintillator element are optically connected to correspond to one body 1. The photodiode has an active area (sensitive area), and converts light received in the active area into an electric signal.

  The switching chip 3285 includes a plurality of CMOS (Complementary MOS) transistors. The plurality of transistors are connected to the plurality of photodiodes. The electrical signal generated by the photodiode is supplied to the DAS chip 3289 via the corresponding transistor.

  The multilayer wiring board 3287 is composed of a plurality of thin film printed wirings provided with via holes. At least one of the thin film printed wirings has an X-ray shielding function.

  The DAS chip 3289 performs amplification processing and A / D conversion processing on the electrical signal sent via the switching chip 3285.

  Instead of providing an X-ray shielding function to at least one of the plurality of thin film printed wirings constituting the multilayer wiring board 3287, as shown in FIG. 26 (b), the DAS chip 3289 is placed outside the X-ray irradiation region, that is, You may make it arrange | position to the edge of the multilayer wiring board 3287. FIG. Further, a lead shielding plate 3271 is disposed above the DAS chip 3289. The lead plate 3271 is attached to a support 3270 for securing the detector module 3280 in the shield housing.

  Further, the shielding plate 3271 may be disposed between the multilayer wiring board 3287 and the DAS chip 3289 as shown in FIG. In this case, the electrical connection between the multilayer wiring board 3287 and the DAS chip 3289 is performed by wire bonding, not by the pull-chip connection.

  As shown in FIG. 26 (d), when a separator 3273 made of lead, molybdenum or the like is inserted between adjacent scintillator elements, X-rays hardly pass through the scintillator 3281, so that the DAS chip 3289 It may be arranged inside the X-ray irradiation area.

  According to such a configuration, X-ray incidence on the DAS chip 3289 can be prevented, and the influence on components having poor X-ray resistance can be prevented. In particular, as shown in FIGS. 26 (c) and 26 (d), if a structure in which lead is sandwiched between the multilayer wiring board 3287 and the DAS chip 3289, or a non-X-ray transmission type separator is adopted, the DAS Since there is no restriction on the arrangement of the chip 3289, the radiation detector can be easily manufactured.

  In manufacturing the detector module 3280, the DAS chip 3289 and the switching chip 3285 are connected to the multilayer wiring board 3287 at 250 ° C. by solder reflow. In the subsequent process, the photodiode chip 3283 is connected to the switching chip 3285 at a relatively low temperature of 120 to 150 ° C. By this procedure, it is possible to avoid the problem that the surface of the photodiode chip 3283 is exposed to the solder atmosphere and becomes dirty.

  In the connection process described above, operating heat generated in the switching chip 3285 and the DAS chip 3289 may be used. This is because the amount of heat generated is about 100 to 200 W for the radiation detector 3028 as a whole, and a sufficient amount of heat for the connection can be secured. Therefore, it is not necessary to use a heater as in the prior art, and no large-scale equipment is required for connection.

  Next, on / off control of the transistors of the switching chip 3285 for changing the slice thickness will be described.

  As shown in FIG. 28, in order to obtain 32 slices of 0.5 mm thickness, 32 transistors connected to 32 0.5 mm wide photodiodes 3001 arranged in the center are serially turned on. . As a result, the electrical signals detected by the 32 photodiodes 3001 are individually supplied to the DAS chip 3289. Charges generated in other photodiodes 3002 are leaked to the ground.

  In addition, as shown in FIG. 29, to obtain 32 1 mm-thick slices can be realized by simultaneously turning on two transistors corresponding to adjacent pairs of photodiodes 3001. That is, 16 transistors connected to the 1 mm wide photodiode 3002 and 16 sets of transistors connected to the pair of photodiodes 3001 are turned on serially.

  As shown in FIG. 30A, an Al wiring 3292 formed on the surface is connected to a photodiode 3294 formed on the surface of the silicon substrate 3285 of the photodiode chip 3283, and this Al wiring 3292 is connected to the silicon substrate. A back surface of the wiring 3299 penetrating through 3285 is led out to the back surface, and a bump 3230 is formed at the tip. By using the through wiring 3299, the wiring can be drawn out to the back surface of the photodiode chip 3283.

  As shown in FIG. 30B, a through wiring 3299 may be formed immediately below the photodiode 3284. In addition, as shown in FIG. 30C, the photodiode chip 3283 and the switching chip 3285 may be micro-bump connected by embedded wiring.

  Note that the photodiode chip 3283 and the switching chip 3285 may be connected by a multilayer wiring board.

  Next, the operation of the detector module 3280 configured as described above will be described.

  Conventionally, since the detector and the data collection device are separate, an analog signal is transmitted from the photodiode to the data collection device 3085 by a flexible PC board. This structure requires a long flexible PC board. Therefore, it is weak against vibration due to rotation at the time of scanning, causing the connector to be pulled out by the flexible PC board and increasing noise. An increase in the number of detection elements leads to an increase in the number of flexible PC boards and further noise generation.

  In contrast, the detector module 3280 has a scintillator 3281, a photodiode chip 3283, a switching chip 3285, and a DAS chip 3289 mounted in common on a rigid multilayer wiring board 3287. Therefore, it is not necessary to take out a signal by a connector or a long flexible PC board.

  Further, a signal taken out from the detector module 3280 and sent to the system unit 3014 is a digital signal. Therefore, problems such as vibration due to rotation during scanning, noise generation due to the antenna effect of the flexible PC board, and disconnection of the connector do not occur. As a result, noise generation and the like can be prevented, and the image quality can be improved.

  Moreover, the signal extraction wiring from a detector increases with many arrangement | positioning of an element. In the signal extraction by the connector and the flexible PC board used in the conventional detector, the apparatus becomes quite large. In addition, the switch chip used in the conventional detector is installed in the same plane as the photodiode. Therefore, the density of signals input to the switch chip is increased, and the active area of the photodiode is reduced. As a result, the sensitivity to X-rays that have passed through the subject is low, and the resulting image is noisy. This noise generation becomes more remarkable when the number of detection elements increases.

  On the other hand, the detector module 3280 has a configuration in which a scintillator 3281, a photodiode chip 3283, a switching chip 3285, a multilayer wiring board 3287, and a DAS chip 3289 are stacked and integrated, so that space can be saved. Can do. The photodiode of the photodiode chip 3283 and the transistor of the switching chip 3285 are electrically connected by bump connection. Therefore, a space-saving electrical wiring is realized in the laminated structure. In addition, this configuration with stacked layers and bump connections can reduce the signal density compared to a conventional detector that inputs a large number of signal wires to a transistor installed in the same plane as the photodiode. A wide area can be secured.

  According to the above configuration, the following effects can be obtained.

(1) High-definition (high-resolution) images can be taken over a wide range by unit time without increasing the size of the apparatus. In addition, the sensitivity to X-rays can be improved, and the quality of the obtained image can be improved.

(2) A flexible PC board and a backplane substrate are not required as in a conventional detector. As a result, the outer shape of the detector unit can be reduced. Furthermore, a DAS chassis is not required, and space saving in the CT apparatus can be realized.

(3) The protrusion of the front cover that has existed in the conventional detector can be eliminated. Therefore, the dome opening can be felt wider than before, and the intimidation of the patient can be reduced. In addition, the accessibility of the operator or doctor to the patient can be improved.

(3) The interference between the gantry and the bed is reduced, and the tilt angle can be deepened. Therefore, it is possible to perform imaging while the patient is in a comfortable posture.

(4) When the detector system according to the present invention is applied to a conventional apparatus, a tomographic image having a slice thickness wider than that of the conventional one can be obtained.

(5) The accessibility to the unit mounted in the back of the detector unit can be improved.

(6) The weight of the detector unit can be reduced. As a result, when the detector unit is rotated during imaging, the influence of shaking of the gantry due to the rotation can be reduced compared to the conventional case.

(7) The number of parts used is smaller than that of the conventional detector, and the cost can be reduced.

(Modification)
The present invention is not limited to the above-described embodiments, and various modifications can be made without departing from the scope of the invention at the stage of implementation. Furthermore, the above embodiment includes various stages, and various inventions can be extracted by appropriately combining a plurality of disclosed constituent elements. For example, some constituent requirements may be deleted from all the constituent requirements shown in the embodiment.

The perspective view of the radiation detector by 1st Embodiment of this invention. The lineblock diagram for the line of the radiation detector by a 1st embodiment of the present invention. Sectional drawing of the radiation detector by 2nd Embodiment. Sectional drawing which shows the connection of the photodiode chip | tip of FIG. 3, and a rigid printed wiring board. Sectional drawing of the radiation detector by 3rd Embodiment. Sectional drawing of the multilayer wiring board of FIG. Sectional drawing of the radiation detector by 4th Embodiment. Sectional drawing of the radiation detector by 5th Embodiment. Sectional drawing of the radiation detector by 6th Embodiment. The block diagram of the X-ray CT apparatus by 7th Embodiment. FIG. 11 is a plan view of one detector module of FIG. 10. The figure which shows the flow of the data processing in the X-ray CT apparatus shown in FIG. The figure which shows the some photodiode selected in accordance with the conditions of slice thickness 0.5mmx16 slice in 7th Embodiment. The figure which shows the some photodiode selected according to the conditions of slice thickness 1mm * 16 slice in 7th Embodiment. The figure which shows the some photodiode selected according to the conditions of slice thickness 1mm * 32 slice in 7th Embodiment. The figure which shows the example of a display of an axial tomogram, a sagittal tomogram, and a coronal tomogram in 7th Embodiment. The figure which shows the example of a display of a moving image in 7th Embodiment. The block diagram of the X-ray CT apparatus by 8th Embodiment. The top view of each of the several detector module which comprises the radiation detector of FIG. (A) is a perspective view of a detector block constituting one detector module of FIG. 19, (b) is a perspective view of a DAS block constituting one detector module of FIG. 19, and (c) is FIG. 20 is a perspective view of the detector module of FIG. 19. (A) is sectional drawing of the detector module of 8th Embodiment, (b) is sectional drawing of the other detector module of 8th Embodiment. Sectional drawing of the DAS block of Fig.21 (a). (A) is sectional drawing of the detector module of 8th Embodiment, (b) is sectional drawing of the other detector module of 8th Embodiment. (A) is a surface view of the detector module of 8th Embodiment, (b) B is a rear view of the detector module of 8th Embodiment. The block diagram of the X-ray CT apparatus of 9th Embodiment. (A) is sectional drawing of the X-ray detector module of 9th Embodiment, (b) is sectional drawing of the other X-ray detector module of 9th Embodiment, (c) is 9th Embodiment. Sectional drawing of other X-ray detector module, (d) D is sectional drawing of the further another X-ray detector module of 9th Embodiment. The top view of the X-ray detector module of 9th Embodiment. The figure which shows signal read-out control corresponding to the conditions of slice thickness of 0.5 mm x 32 slices in 9th Embodiment. The figure which shows signal read-out control corresponding to the conditions of slice thickness 1mm * 32 slice in 9th Embodiment. (A) is schematic sectional drawing of a switching chip in 9th Embodiment, (b) is schematic sectional drawing of another switching chip in 9th Embodiment, (c) is further in 9th Embodiment, The schematic sectional drawing of another switching chip.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... Detector module, 2 ... Photodiode chip, 3 ... Photodiode 3, 4 ... Switching chip, 5 ... Transistor, 7 ... Signal read-out line, 8 ... DAS chip, 9 ... Scintillator 9, 10 ... Rigid printed wiring board, 14: Al wiring.

Claims (6)

A scintillator that converts X-rays incident from the surface side into light;
At least one photodiode chip comprising a plurality of photodiodes for converting the converted light into an electrical signal;
At least one switching chip comprising a plurality of switching elements for reading a plurality of signals from the plurality of photodiodes;
At least one data collection chip comprising a plurality of data collection units for amplifying and digitizing the plurality of read signals;
A rigid multilayer wiring board that mounts the photodiode chip, the switching chip, and the data collection chip in common;
The back surface of the switching chip is coupled to the surface of the multilayer wiring board, the back surface of the photodiode chip is coupled to the surface of the switching chip, and the data collection chip is coupled to the back surface of the multilayer wiring board. Radiation detection system. The plurality of photodiodes are connected to the plurality of switching elements via a plurality of first through wirings penetrating the photodiode substrate from the front surface to the back surface.
The plurality of switching elements are connected to a plurality of surface wirings of the multilayer wiring board via a plurality of second through wirings penetrating the switching element substrate from the front surface to the back surface,
Wherein the data acquisition chip, the radiation detection system according to claim 1, characterized in that flip-chip connected to a plurality of backside interconnect of the multilayer wiring board. Wherein the data acquisition chip, the radiation detection system according to claim 1, characterized in that it is disposed at the periphery of the multilayer wiring board. The radiation detection system according to claim 3 , further comprising a plurality of X-ray shielding plates disposed above the data collection chip. A scintillator that converts X-rays incident from the surface side into light;
At least one photodiode chip comprising a plurality of photodiodes for converting the converted light into an electrical signal;
At least one switching chip comprising a plurality of switching elements for reading a plurality of signals from the plurality of photodiodes;
At least one data collection chip comprising a plurality of data collection units for amplifying and digitizing the plurality of read signals;
A first rigid printed wiring board for mounting the photodiode chip and the switching chip;
A second rigid printed wiring board on which the data collection chip is mounted;
A radiation detection system comprising: a connector for detachably connecting the first rigid printed wiring board and the second rigid printed wiring board. The photodiode chip is bump-connected to the surface wiring of the first rigid printed wiring board via a plurality of through wirings that penetrate the semiconductor substrate from the front surface to the back surface and are connected to the plurality of photodiodes. The radiation detection system according to claim 5 .

Images