U.S. patent application number 11/102795 was filed with the patent office on 2005-08-18 for semiconductor radiation detector, positron emission tomography apparatus, semiconductor radiation detection apparatus, detector unit and nuclear medicine diagnostic apparatus.
Invention is credited to Amemiya, Kensuke, Kitaguchi, Hiroshi, Kojima, Shinichi, Tsuchiya, Katsutoshi, Ueno, Yuuichirou, Yanagita, Norihito, Yokoi, Kazuma, Yokomizo, Osamu.
Application Number | 20050178969 11/102795 |
Document ID | / |
Family ID | 34309054 |
Filed Date | 2005-08-18 |
United States Patent
Application |
20050178969 |
Kind Code |
A1 |
Amemiya, Kensuke ; et
al. |
August 18, 2005 |
Semiconductor radiation detector, positron emission tomography
apparatus, semiconductor radiation detection apparatus, detector
unit and nuclear medicine diagnostic apparatus
Abstract
Each semiconductor radiation detector used for a nuclear
medicine diagnostic apparatus (PET apparatus) is constructed with
an anode electrode A facing a cathode electrode C sandwiching a
CdTe semiconductor member S which generates charge through
interaction with .gamma.-rays. Then, a thickness t of the
semiconductor member S sandwiched between these mutually facing
anode electrode A and cathode electrode C is set to 0.2 to 2 mm.
Furthermore, the devices are mounted (laid out) on substrates in
such a way that the distance (distance of conductor) between the
semiconductor radiation detector and an analog ASIC which processes
the signal detected by this detector is shortened. Furthermore, the
substrates on which the detectors are mounted are housed in a
housing as a unit (detector unit).
Inventors: |
Amemiya, Kensuke;
(Hitachinaka, JP) ; Ueno, Yuuichirou; (Hitachi,
JP) ; Kitaguchi, Hiroshi; (Naka, JP) ;
Yokomizo, Osamu; (Tokai, JP) ; Kojima, Shinichi;
(Hitachi, JP) ; Tsuchiya, Katsutoshi; (Hitachi,
JP) ; Yanagita, Norihito; (Hitachi, JP) ;
Yokoi, Kazuma; (Hitachi, JP) |
Correspondence
Address: |
DICKSTEIN SHAPIRO MORIN & OSHINSKY LLP
2101 L Street, NW
Washington
DC
20037
US
|
Family ID: |
34309054 |
Appl. No.: |
11/102795 |
Filed: |
April 11, 2005 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
11102795 |
Apr 11, 2005 |
|
|
|
10874343 |
Jun 24, 2004 |
|
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Current U.S.
Class: |
250/363.05 |
Current CPC
Class: |
G01T 1/2928 20130101;
G01T 1/2985 20130101 |
Class at
Publication: |
250/363.05 |
International
Class: |
G01T 001/166 |
Foreign Application Data
Date |
Code |
Application Number |
Sep 30, 2003 |
JP |
2003-340688 |
Claims
1-55. (canceled)
56. A nuclear medicine diagnostic apparatus, comprising: a support
member; and a plurality of detector units detachably mounted on
said support member, said detector units having a housing member
and a plurality of unit substrates detachably placed in said
housing member, wherein said unit substrates include a plurality of
semiconductor radiation detectors to which radiation is input and
an integrated circuit for processing radiation detection signals
output from each of said semiconductor radiation detectors.
57. The nuclear medicine diagnostic apparatus according to claim
56, further comprising a bed for laying an object to be examined,
wherein said plurality of detector units are arranged around said
bed, and said semiconductor radiation detectors are placed near
said bed and said integrated circuit is placed far from said bed in
said unit substrates.
58. The nuclear medicine diagnostic apparatus according to claim
56, further comprising a tomographic information creation apparatus
for creating tomographic information using second information
obtained based on first information output from said integrated
circuit.
59. A nuclear medicine diagnostic apparatus, comprising: a support
member; and a plurality of detector units detachably mounted on
said support member, said detector units having a housing member
and a plurality of unit substrates detachably placed in said
housing member, wherein said unit substrates include a plurality of
semiconductor radiation detectors to which radiation is input, an
integrated circuit for processing radiation detection signals
output from each of said semiconductor radiation detectors and a
counting apparatus to which outputs from said integrated circuit
are input for counting in pairs said outputs corresponding to a
pair of radiation detection signals detected within a set time.
60. The nuclear medicine diagnostic apparatus according to claim
59, further comprising a tomographic information creation apparatus
for creating tomographic information using counting information
output from said counting apparatus.
61. The nuclear medicine diagnostic apparatus according to claim
56, wherein said unit substrates comprises a first substrate
including said semiconductor radiation detectors and a second
substrate including said integrated circuit.
62. The nuclear medicine diagnostic apparatus according to claim
56, wherein said integrated circuit comprises: an analog integrated
circuit for processing signals output from said semiconductor
radiation detectors; an AD converter for converting analog signals
output from said analog integrated circuit to digital signals; and
a digital integrated circuit to which signals from said AD
converter are input.
63. The nuclear medicine diagnostic apparatus according to claim
56, wherein said housing member comprises guide members for guiding
said unit substrates into said housing member.
64. The nuclear medicine diagnostic apparatus according to claim
56, wherein said housing member has light shielding properties.
65. The nuclear medicine diagnostic apparatus according to claim
62, wherein said detector units comprise a separate integrated
circuit to which information output from said digital integrated
circuit is input.
66. A positron emission tomography apparatus, comprising: a support
member; and a plurality of detector units detachably mounted on
said support member, said detector units having a housing member
and a plurality of unit substrates detachably placed in said
housing member, wherein said unit substrates include a plurality of
semiconductor radiation detectors to which radiation is input and
an integrated circuit for processing radiation detection signals
output from each of said semiconductor radiation detectors, wherein
said integrated circuit outputs time information of radiation
detection and identification information of said semiconductor
radiation detectors which detected radiation.
67. The positron emission tomography apparatus according to claim
66, wherein said integrated circuit comprises: an analog integrated
circuit for processing signals output from said semiconductor
radiation detectors; an AD converter for converting analog signals
output from said analog integrated circuit to digital signals; and
a digital integrated circuit to which signals from said AD
converter are input.
68. A detector unit, comprising: a unit substrate including a
plurality of semiconductor radiation detectors and an integrated
circuit for processing signals output from said semiconductor
radiation detectors, wherein said plurality of semiconductor
radiation detectors comprises a plurality of signal lines for
transmitting output signals from each of said semiconductor
radiation detectors to said integrated circuit; and a housing
member for housing said plurality of semiconductor radiation
detectors.
69. The detector unit according to claim 68, wherein said housing
member has light shielding properties.
70. The detector unit according to claim 68, wherein said unit
substrates comprise a first substrate including said semiconductor
radiation detectors and a second substrate including said
integrated circuit.
71. The detector unit according to claim 68, wherein said
integrated circuit comprises: an analog integrated circuit for
processing signals output from said semiconductor radiation
detectors; an AD converter for converting analog signals output
from said analog integrated circuit to digital signals; and a
digital integrated circuit to which signals from said AD converter
are input.
72. The detector unit according to claim 70, wherein said first and
second substrates are overlapped at end portions thereof and
detachably connected with each other.
73. The detector unit according to claim 68, wherein said
semiconductor radiation detectors are arranged on both planes of
said unit substrates.
74. The detector unit according to claim 70, wherein said
semiconductor radiation detectors are arranged on both planes of
said first substrate.
75. The detector unit according to claim 71, further comprising a
separate integrated circuit placed in said housing member, to which
information output from said digital integrated circuit is input.
Description
CROSS-REFERENCE TO RELATED APPLICATION
[0001] The present application is related to a U.S. Ser. No. ______
being filed based on Japanese Patent Application No. 2003-342437
filed on Sep. 30, 2003, the entire content of which is incorporated
herein by reference.
BACKGROUND OF THE INVENTION
[0002] The present invention relates to a nuclear medicine
diagnostic apparatus, and more particularly, to a positron emission
tomography (hereinafter referred to as "PET") apparatus, which is a
kind of a nuclear medicine diagnostic apparatus using a
semiconductor radiation detector, semiconductor radiation detection
apparatus or detector unit.
[0003] A detector using a NaI scintillator is known as a
conventional radiation detector for detecting radiation such as
.gamma.-rays. With a gamma camera (a kind of nuclear medicine
diagnostic apparatus) provided with a NaI scintillator, radiation
(.gamma.-rays) incident on the scintillator at an angle restricted
by many collimators interacts with NaI crystals and emits
scintillation light. This light travels in such a way as to
sandwich a light guide, reaches a photoelectric multiplier and
becomes an electrical signal. The electrical signal is shaped by a
measuring circuit mounted on a measuring circuit fixing board and
transferred from an output connector to an external data collection
system. All these scintillator, light guide, photoelectric
multiplier and measuring circuit, measuring circuit fixing board,
etc., are housed in a light shielding case and shielded from
electromagnetic waves other than external radiation.
[0004] Since a gamma camera using a scintillator has a structure
with a large photoelectric multiplier (also called
"photomultiplier") placed after one large crystal such as NaI, its
position resolution remains on the order of 10 mm. Furthermore,
since the scintillator detects radiation in multi-stages of
conversion from radiation to visible light, from visible light to
electrons, it has a problem of having considerably poor energy
resolution. For example, there is a PET apparatus (positron
emission tomography apparatus) having position resolution of 5 to 6
mm or a high-end PET apparatus having position resolution of 4 mm
or so, but since their photoelectric multipliers use vacuum tubes,
it is difficult to further improve position resolution.
[0005] There are radiation detectors for detecting radiation
according to principles different from those of such a
scintillator, such as semiconductor radiation detectors provided
with a semiconductor radiation detection element using a
semiconductor material such as CdTe (cadmium telluride), TlBr
(thallium bromide) and GaAs (gallium arsenide).
[0006] This semiconductor radiation detector is attracting
attention because its semiconductor radiation detection element
converts electrical charge produced by interaction between
radiation and the semiconductor material to an electrical signal,
and therefore it has better efficiency of conversion to an
electrical signal than the scintillator and can also be
miniaturized.
[0007] [Patent Document 1] JP-A-2003-79614 (paragraph No. 0016)
[0008] [Patent Document 2] JP-A-2003-167058 (paragraph No. 0020,
0023)
[0009] Meanwhile, when a semiconductor material such as Tl making
up a semiconductor radiation detection element interacts with
radiation in a semiconductor radiation detector, holes having
positive electrical charge and electrons having negative electrical
charge are generated. While mobility of electrons is relatively
large, mobility of holes is relatively small. That is, electrons
move relatively easily and holes move with difficulty. This takes
more time for holes to reach an electrode than electrons. Moreover,
holes may be annihilated before reaching the electrode. This
involves a problem that the detection sensitivity of radiation is
worsened. Thus, these problems require solutions.
[0010] It is an object of the present invention to provide a
semiconductor radiation detector capable of improving detection
sensitivity.
SUMMARY OF THE INVENTION
[0011] In order to solve the above described problems, a first
embodiment of the present invention improves detection sensitivity
by shortening a distance between electrodes for charge collection
of a semiconductor radiation detector. That is, the distance
between an anode electrode and cathode electrode or the thickness
of a semiconductor area sandwiched between the anode electrode and
cathode electrode is 0.2 to 2 mm. In this structure, the distance
from positions of electrons and holes generated by interaction
between the semiconductor material and radiation to the electrodes
is shortened, and therefore the time required for them to reach the
electrodes is shortened. Furthermore, shortening the distance up to
the electrodes reduces the probability that holes may be
annihilated midway the distance.
[0012] A second embodiment of the present invention is a nuclear
medicine diagnostic apparatus comprising a plurality of unit
substrates including a plurality of semiconductor radiation
detectors for introducing radiation and an integrated circuit for
processing radiation detection signals output from the plurality of
semiconductor radiation detectors. This allows the semiconductor
radiation detectors and the integrated circuits which process the
outputs to be disposed close to one another, with the result that
when weak output signals of the semiconductor radiation detectors
are transmitted to the integrated circuits, it is possible to
reduce influences of noise on the weak output signals.
[0013] The semiconductor radiation detector, analog LSI (Large
Scale Integrated Circuit), AD converter and digital LSI are
preferably arranged on the unit substrate in that order and the
respective elements are connected by wiring so that a signal
detected by the semiconductor radiation detector is processed by
the analog LSI, the signal processed by the analog LSI is processed
by the AD converter, and the signal processed by the AD converter
is processed by the digital LSI. By shortening the distance between
the semiconductor radiation detector and analog LSI in particular,
this structure can shorten the wiring distance between the
semiconductor radiation detector and analog LSI and thereby reduce
noise superimposed on the wiring until the signal detected by the
semiconductor radiation detector reaches the analog LSI. In an
embodiment which will be described later, the LSI (integrated
circuit) corresponds to an ASIC. Also, the semiconductor radiation
detection apparatus corresponds to a combined substrate (detector
substrate+ASIC substrate) in the embodiment which will be described
later.
[0014] According to the second embodiment, detection signals when
the semiconductor radiation detectors detect radiation are
processed by an application-specific IC called "ASIC (Application
Specific Integrated Circuit)" and this embodiment is intended to
solve an additional problem discovered by the inventor et al. that
since the detection signals output from the semiconductor radiation
detectors are weak, the ASIC is easily affected by noise. A
reduction of the noise leads to substantial improvement of
detection sensitivity (count, peak value, time detection accuracy)
by the semiconductor radiation detectors.
[0015] Different substrates are preferably used as the substrate
for mounting the semiconductor radiation detectors and the
substrate for mounting the LSI. During ordinary operation, the two
substrates are used in combination as a combined substrate (unified
substrate) so that in the event of trouble, only the troubled
substrate can be replaced to thereby facilitate maintenance and
examination, etc.
[0016] A third embodiment of the present invention adopts a
unit-type construction in which a plurality of unit substrates
including semiconductor radiation detectors and an integrated
circuit are mounted in a frame in a detachable/attachable manner.
Since it is only necessary to mount a detector unit including a
plurality of unit substrates on a nuclear medicine diagnostic
apparatus, a plurality of semiconductor radiation detectors can be
mounted on the nuclear medicine diagnostic apparatus at a time. In
this way, the time required to mount the semiconductor radiation
detectors on the nuclear medicine diagnostic apparatus can be
shortened drastically.
[0017] The embodiment is preferably adapted so that these unit
substrates can be removed from the detector unit one by one or the
whole detector unit can be removed from the nuclear medicine
diagnostic apparatus, or more specifically, from the camera, which
facilitates maintenance and examination.
[0018] Note that many semiconductor radiation detectors are used
for a nuclear medicine diagnostic apparatus (radiological
diagnostic apparatus) such as PET, SPECT and gamma camera. For
example, a PET uses a hundred thousand to several hundreds of
thousands of (channels) semiconductor radiation detectors and there
is a demand for shortening the time required to mount these many
semiconductor radiation detectors on the nuclear medicine
diagnostic apparatus. A fourth embodiment of the present invention
is implemented to meet such a demand. There is also a demand for
facilitating maintenance and examination of semiconductor radiation
detectors.
[0019] The present invention can prevent or reduce deterioration of
the detection sensitivity of radiation using semiconductor
radiation detectors. The present invention can also prevent or
reduce deterioration of signals detected by the semiconductor
radiation detectors. This allows, for example, a nuclear medicine
diagnostic apparatus to obtain clear images.
[0020] Other objects, features and advantages of the invention will
become apparent from the following description of the embodiments
of the invention taken in conjunction with the accompanying
drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
[0021] FIG. 1 is a perspective view showing a structure of a PET
apparatus as a nuclear medicine diagnostic apparatus according to
this embodiment;
[0022] FIG. 2 schematically shows a cross section in a
circumferential direction of the camera of the PET apparatus in
FIG. 1;
[0023] FIG. 3 schematically shows a structure of a semiconductor
radiation detector in a minimum construction;
[0024] FIG. 4 is a graph comparing a "time-peak value curve"
between a case where a thickness t of a semiconductor material of
the semiconductor radiation detector is large and a case where it
is small;
[0025] FIG. 5 is a graph schematically showing a relationship
between the thickness t of a semiconductor material and a peak
value (maximum value) of a semiconductor radiation detector;
[0026] FIG. 6 schematically illustrates a construction of a
semiconductor radiation detector having a laminated structure of
semiconductor materials and electrodes (anodes, cathodes);
[0027] FIG. 7A is a front view of a combined substrate which
combines a detector substrate and an ASIC substrate of the
semiconductor radiation detectors according to this embodiment,
FIG. 7B is a side view of FIG. 7A and FIG. 7C is a perspective view
schematically showing a construction of the semiconductor radiation
detector mounted on the detector substrate in FIG. 7A;
[0028] FIG. 8 is a block diagram schematically showing an analog
detection circuit;
[0029] FIG. 9 is a block diagram showing a schematic construction
of a digital ASIC and a connection relationship between an analog
ASIC and the digital ASIC;
[0030] FIG. 10 is a perspective view quoted to illustrate a
construction of a detector unit housing a plurality of
semiconductor radiation detectors;
[0031] FIG. 11 is a side view of the detector unit in FIG. 10
without the side plate;
[0032] FIG. 12A is a partially exploded perspective view of a
camera when the detector unit is mounted on the camera and FIG. 12B
is a cross-sectional view of the central part of the camera;
[0033] FIG. 13 is a perspective view showing a construction of a
SPECT apparatus as a nuclear medicine diagnostic apparatus
according to another embodiment;
[0034] FIG. 14 is a block diagram schematically showing a circuit
construction of an analog ASIC of the SPECT apparatus in FIG. 13;
and
[0035] FIG. 15 is a block diagram showing a schematic construction
of a digital ASIC in the SPECT apparatus in FIG. 13 and a
connection relationship between the analog ASIC and digital
ASIC.
DESCRIPTION OF THE EMBODIMENTS
Embodiment 1
[0036] A nuclear medicine diagnostic apparatus which is a preferred
embodiment of the present invention will be explained with
reference to attached drawings in detail below as appropriate. The
following are explanations of the nuclear medicine diagnostic
apparatus according to this embodiment, distance between electrodes
of a semiconductor radiation detector, arrangement (layout) of
elements such as analog ASIC on a substrate, and elements
applicable to this embodiment for construction of substrate units,
etc. Note that an analog ASIC refers to an ASIC (Application
Specific Integrated Circuit) which is an application-specific IC
for processing analog signals and is a kind of LSI.
[0037] <<Nuclear Medicine Diagnostic Apparatus>>
[0038] First, the nuclear medicine diagnostic apparatus
(radiological diagnostic apparatus) according to this embodiment
will be explained. As shown in FIG. 1, a PET apparatus 1 as the
nuclear medicine diagnostic apparatus is constructed by including a
camera (image pickup apparatus) 11, a data processing apparatus 12,
a display apparatus 13, etc. An examinee is laid on a bed 14 to be
photographed using the camera 11. The camera 11 incorporates many
semiconductor radiation detectors 21 (see FIG. 3, FIGS. 7A-7C, FIG.
10) to detect .gamma.-rays emitted from the body of the examinee
using semiconductor radiation detectors (hereinafter simply
referred to as "detectors") 21. The camera 11 is provided with an
integrated circuit (ASIC) for measuring peak values, detection
times of .gamma.-rays and is designed to measure peak values and
detection times of detected radiations (.gamma.-rays). The data
processing apparatus 12 includes a storage apparatus, a
simultaneous measuring apparatus 12A (see FIG. 2) and a tomographic
information creation apparatus 12B (see FIG. 2). The data
processing apparatus 12 takes in data of peak values, detection
times of detected .gamma.-rays and packet data including detector
(channel) IDs. The simultaneous measuring apparatus 12A carries out
simultaneous measurements based on this packet data, especially
data of detection times and detector IDs, identifies detection
positions of 511 KeV .gamma.-rays and stores them in the storage
apparatus. The tomographic information creation apparatus 12B
creates a functional image based on the identified positions and
displays it on the display apparatus 13.
[0039] As shown in FIG. 2, inside the camera 11, many detector
units 2 (see FIG. 10 for details) housing a plurality of combined
substrates 20 (see FIG. 7 for details) provided with many detectors
21 for detecting .gamma.-rays emitted from the examinee are
disposed circumferentially. The examinee is laid on the bed 14 and
positioned at the center of the camera 11. At this time, the
detectors are disposed so as to surround the bed 14. The detector
unit 2 is designed to output for each detector 21 included in the
detector unit 2, peak value information of .gamma.-rays obtained
based on a detection signal when a detector 21 interact with
.gamma.-rays, time information on .gamma.-ray detection and address
information (detector ID) of each detector 21. The constructions of
the detector 21, combined substrate 20 and detector unit 2 will be
explained in detail later. The examinee is administered
radiopharmaceuticals, for example, fluorodeoxyglucose (FDG)
containing .sup.18F whose half-life is 110 minutes. .gamma.-rays
(annihilated .gamma.-rays) are emitted from the body of the
examinee when positrons emitted from the FDG annihilate.
[0040] Hereafter, the characteristic parts of this embodiment will
be explained.
[0041] <<Semiconductor Radiation Detector; Distance Between
Electrodes>>
[0042] First, the detector 21 applied to this embodiment will be
explained. As shown in FIG. 3, the detector 21 is constructed of a
semiconductor radiation detection element (hereinafter referred to
as "detection element") 211 made of a tabular semiconductor
material S, both sides of which are covered with thin-plate (film)
electrodes (anode A, cathode C) (minimum construction). Of these
components, the semiconductor material S is made up of a single
crystal of any one of the above described CdTe (cadmium telluride),
TlBr (thallium bromide), GaAs (gallium arsenide), etc. Furthermore,
the electrodes (anode A, cathode C) are made of any one material of
Pt (platinum), Au (gold), In (indium), etc. In the following
explanations, suppose the semiconductor material S is obtained by
slicing a CdTe single crystal. Furthermore, suppose radiation to be
detected is 511 KeV .gamma.-rays used for the PET apparatus.
[0043] An overview of the principle of .gamma.-ray detection using
the detector 21 will be explained using FIG. 3. When .gamma.-rays
are introduced into the detector 21 and interaction occurs between
.gamma.-rays and the semiconductor material S constituting the
detector 21, an amount of hole and electron pairs schematically
shown in the figure with "+" and "-" corresponding to the energy of
.gamma.-rays is generated. Here, a voltage (e.g., 300 V) for charge
collection is applied between the electrodes of the anode A and
cathode C of the detector 21. Because of this, holes are moved
attracted to the cathode C and electrons are moved attracted to the
anode A. When holes and electrons are compared, as described in
"Disclosure of the invention" the ease of movement (mobility) of
electrons is relatively large and therefore electrons reach the
anode in a shorter time. On the other hand, the mobility of holes
is relatively small and therefore holes take more time to reach the
cathode. Note that holes may be annihilated before reaching the
electrode.
[0044] As shown in FIG. 4 which shows a comparison in the
"time-peak value curve" between a case where the semiconductor
material S (detection element 211) of the detector 21 is thick and
a case where it is thin, the semiconductor material S having a
smaller thickness t has a quicker rise of peak value and a higher
maximum value of the peak value. Having a quicker rise of the peak
value contributes to improvement of the accuracy of simultaneous
measurement of the PET, for example. Furthermore, having a higher
peak value contributes to increasing energy resolution. Thus, a
smaller thickness t speeds up rising of the peak value and
increases the peak value (the efficiency of charge collection
improves) because the time for electrons and holes to reach the
electrodes (anode A, cathode C) (time of charge collection) is
shortened. This is also because holes which would be conventionally
annihilated midway can reach the electrode (cathode C) without
annihilation because of the shorter distance. Note that the
thickness t can also be expressed by the distance between the
electrodes, anode A and cathode C facing each other.
[0045] The thickness (distance between electrodes) t of the
detection element 211 is preferably 0.2 mm to 2 mm. This is because
a thickness t of not less than 2 mm slows down the rising speed of
the peak value and reduces the maximum value of the peak value as
well. On the other hand, a thickness t of smaller than 0.2 mm
relatively increases the thickness (volume) of the electrodes
(anode, cathode) and when installed on a substrate, the proportion
of the very semiconductor material S that interacts with radiation
decreases. That is, reducing the thickness t of the semiconductor
material S relatively increases the thickness of the electrode
which does not interact with .gamma.-rays on one hand, and the
proportion of the semiconductor material S which interacts with
.gamma.-rays relatively decreases on the other, with the result
that the sensitivity of detecting .gamma.-rays decreases
(.gamma.-rays pass by). Furthermore, a smaller thickness t may
cause more leakage current preventing a high voltage from being
applied for charge collection.
[0046] For the same reason, the thickness t of the semiconductor
material S is preferably 0.5 mm to 1.5 mm and such a thickness t
allows more reliable detection of .gamma.-rays and more correct
measurement of the peak value, etc.
[0047] In the case of the PET apparatus 1, since it carries out
simultaneous measurement, one of problems to be solved is to
correctly measure a .gamma.-ray detection time. For example, in
FIG. 3, there is a difference in a detection time when positions at
which .gamma.-rays interact with the semiconductor material S are
closer to the cathode C and when those positions are closer to the
anode A. That is, since the moving speed of holes is lower, the
detection time when the interaction occurs closer to the anode A is
relatively late, while the detection time when the interaction
occurs closer to the cathode C is relatively early (approximates to
a real time). That is, also when .gamma.-rays interact with the
semiconductor material S in the same detection element 211, there
is a problem that the detection time changes depending on the
position at which the interaction takes place. More specifically,
when the thickness t is large, the difference in the detection time
depending on the position at which the interaction takes place
increases. Such an event constitutes no big problem in other
fields, but it constitutes a big problem in the case of the PET
apparatus 1, which carries out simultaneous measurement
(simultaneous counting) on the order of nsec (nanoseconds).
Therefore, in this sense, too, it is possible to determine the
detection time appropriately within the above described thickness
range. The detection time is determined by the PET according to an
LET system or CFD system.
[0048] As shown in FIG. 5 which schematically shows a relationship
between the thickness t of the semiconductor material S and peak
value (maximum value) of the detector 21, the maximum value of the
peak value decreases as the thickness t of the semiconductor
material S increases. One of reasons that the peak value decreases
is that holes are annihilated before reaching the electrode. When
the thickness t becomes 2 mm, the peak value of detected radiation
falls short of a threshold whereby it is possible to discriminate
511 KeV .gamma.-rays, and therefore it is not preferable to
increase the thickness t of the semiconductor material S more than
2 mm as described above.
[0049] As shown in FIG. 6, the detector 21 includes the
semiconductor materials S (detection elements 211) laminated in
five layers each sandwiched between the cathode C and anode A. Each
layer of the semiconductor material S is a single layer detector 21
having the aforementioned thickness t (0.2 to 2 mm (more preferably
0.5 to 1.5 mm)). The thickness of the anode A and cathode C is
approximately 20 microns. In the detector 21 having a laminated
structure shown in this FIG. 6, the different anodes A or different
cathodes C are connected to a common wire, and therefore each layer
is not designed to detect radiation independently of other layers.
In other words, when .gamma.-rays interact with the semiconductor
material S, it is not possible to discriminate whether the
interaction takes place on the top layer or bottom layer. Of
course, it is also possible to adopt a structure in which radiation
is detected by each layer. This five-layer structure is adopted
because reducing the thickness t of the semiconductor material S is
preferable in increasing the rising speed of the peak value and
increasing the maximum value of the peak value, but a small
thickness t causes more .gamma.-rays to pass by, and therefore
reducing the amount of .gamma.-rays that pass by while increasing
the efficiency of charge collection increases interaction between
the semiconductor material S and .gamma.-rays (to increase a count
value).
[0050] Adopting the detector 21 having such a laminated structure
can obtain a better peak value rising speed and an accurate peak
value and increase the number of .gamma.-rays (count value)
(increase the sensitivity) that interact with the semiconductor
material S.
[0051] An area s of the electrode (anode A, cathode C) is
preferably 4 to 120 mm.sup.2. An increase of the area s increases
the capacity (stray capacitance) of the detector 21 and this
increase in the stray capacitance makes noise easier to
superimpose, and therefore the area s of the electrode is
preferably as small as possible. Furthermore, charge produced when
.gamma.-rays are detected is partially accumulated in the stray
capacitance, and therefore there is a problem that when the stray
capacitance increases, the amount of charge stored in a charge
amplifier 24b of an analog ASIC 24 or further an output voltage
(peak value) decreases. When CdTe is used for the detector 21, its
dielectric constant is 11 and if the area s of the detector 21 is
120 mm.sup.2, thickness t is 1 mm, then the capacity is 12 pF,
which is not negligible considering the fact that the stray
capacitance of connectors, etc. of the circuit is several pF.
Therefore, the area s of the electrode is preferably 120 mm.sup.2
or less.
[0052] Furthermore, the lower limit of the area s of the electrode
is determined by position resolution of the PET apparatus. The
position resolution of the PET apparatus is determined by not only
the size (array pitch) of the detector 21 but also the positron
range, etc., but since the range of positron of 18F is 2 mm,
setting the size of the detector 21 to 2 mm or less is meaningless.
The method of mounting so that the area of the electrode becomes a
minimum is a case where the surface of the electrode is placed
perpendicular to the radius direction of the camera 11 and from the
above described consideration, the lower limit of one side of the
electrode is 2 mm and the lower limit of the area s of the
electrode is 4 mm.sup.2.
[0053] In the above described explanations, CdTe is used as the
semiconductor material S which interacts with .gamma.-rays, but it
goes without saying that the semiconductor material S may also be
TlBr or GaAs, etc. Moreover, the terms "laminated structure",
"upper layer" and "lower layer" have been used, but these terms are
based on FIG. 6 and when the viewing direction is turned by
90.degree. toward the horizontal direction, the laminated structure
may be read as a parallel structure and top/bottom may be read as
right/left, for example. Moreover, the direction of incident
.gamma.-rays may also be upward, downward, rightward and leftward
in FIG. 6. In other words, the detector 21 has a structure in which
a plurality of (e.g., five) semiconductor materials S are arranged
in parallel in such a way as to sandwich cathodes C and anodes A
alternately.
[0054] <<Combined Substrate; Detector Substrate and ASIC
Substrate>>
[0055] A detailed structure of the combined substrate (unit
substrate) 20 installed in the detector unit 2 (FIG. 10) will be
explained using FIGS. 7A-7C. The combined substrate (semiconductor
radiation detection apparatus) 20 comprises a detector substrate
(first substrate) 20A in which a plurality of detectors 21 are
arranged and an ASIC substrate (second substrate) 20B in which a
capacitor 22, a resistor 23, analog ASICs 24, analog/digital
converters (hereinafter referred to as "ADC") 25 and a digital ASIC
26 are arranged.
[0056] (Detector Substrate)
[0057] With reference to FIGS. 7A-7C, the detector substrate 20A
provided with the detectors 21 will be explained. As shown in FIG.
7A, the detector substrate 20A has a grid-like arrangement
(mounting) of a plurality of detectors 21 on one side of a
substrate body 20a (4 rows of 16 detectors 21=horizontal
16.times.vertical 4=total 64 detectors). In the radius direction of
the camera 11, the detectors 21 are arranged in four rows on the
substrate body 20a. The 16 detectors 21 in the horizontal direction
are arranged in the axial direction of the camera 11, that is, in
the longitudinal direction of the bed 14. Furthermore, as shown in
FIG. 7B, since the semiconductor radiation detectors 21 are
arranged on both sides of the detector substrate 20A, a total of
128 detectors 21 are arranged on one detector substrate 20A. Here,
as the number of detectors 21 to be installed increases, it is
easier to detect .gamma.-rays and it is possible to increase
position accuracy when .gamma.-rays are detected. For this reason,
the detectors 21 are disposed on the detector substrate 20A as
densely as possible. In FIG. 7A, when .gamma.-rays emitted from the
examinee on the bed 14 travel from bottom to top in the figure
(direction indicated by an arrow 32, that is, radius direction of
the camera 11), arranging the detectors 21 in the left-to-right
direction densely on the detector substrate 20A is preferable
because in this way, the number of .gamma.-rays that pass by (the
number of .gamma.-rays that pass through the gap between the
detectors 21) is reduced. This increases the detection efficiency
of .gamma.-rays and increases spatial resolution of an image
obtained.
[0058] As shown in FIG. 7B, the detector substrate 20A of this
embodiment arranges the detectors 21 on both sides of the substrate
body 20a, and therefore the substrate body 20a can be shared by
both sides compared to the case where the detectors 21 are arranged
only on one side. For this reason, it is possible to reduce the
number of substrate bodies 20a by half and arrange the detectors 21
more densely in the circumferential direction. Moreover, as
described above, the number of detector substrates 20A (combined
substrates 20) can be reduced by half, and therefore there is a
merit of saving time and trouble to mount the combined substrates
20 in the housing 30 (see FIG. 10) which will be described
later.
[0059] In the above described explanations, the 16 horizontal
detectors 21 are arranged in the axial direction of the camera 11,
but the arrangement is not limited to this. For example, the 16
horizontal detectors 21 may also be arranged in the circumferential
direction of the camera 11.
[0060] As shown in FIG. 7C, each detector 21 has a laminated
structure of single crystals of the aforementioned thin-film
semiconductor materials S (detection elements 211). The structure
and function thereof have already been explained with reference to
FIG. 6, but supplementary explanations will be given here. As
described above, the detector 21 is provided with the anode A and
cathode C and a potential difference (voltage) of, for example, 300
V is applied between the anode A and cathode C for charge
collection. This voltage is supplied from the ASIC substrate 20B to
the detector substrate 20A via the connector C1 (FIG. 7A).
Furthermore, the signal detected by each detector 21 is supplied to
the ASIC substrate 20B via the connector C1. Thus, on-board wiring
(for charge collection and signal exchange) (not shown) for
connecting the connector C1 and each detector 21 is provided in the
substrate body 20a of the detector substrate 20A. This on-board
wiring has a multi-layered structure. In this embodiment, the
detection elements 211 of the detector 21 are arranged in parallel
to the substrate body 20a. However, the detectors 21 may also be
provided so that the respective detection elements 211 are disposed
perpendicular to the substrate body 20a.
[0061] (ASIC Substrate)
[0062] Then, the ASIC substrate 20B incorporating the ASIC will be
explained. As shown in FIG. 7A, the ASIC substrate 20B is provided
with two analog ASICs 24 and one digital ASIC 26 on one side of the
substrate body 20b. Furthermore, as shown in FIG. 7B, since the
analog ASICs 24 are provided on both sides of the substrate body
20b, one ASIC substrate 20B includes a total of four analog ASICs
24. Furthermore, the ASIC substrate 20B includes eight (=4.times.2)
ADCs 25 on one side of the substrate body 20b and sixteen ADCs 25
on both sides. Furthermore, as many capacitors 22 and resistors 23
as the detectors 21 are arranged on both sides of one substrate
body 20b. Furthermore, to electrically connect theses capacitors
22, resistors 23, analog ASICs 24, ADCs 25 and digital ASIC 26, the
ASIC substrate 20B (substrate body 20b) is provided with on-board
wiring (not shown) as with the above described detector substrate
20A. This on-board wiring also has a laminated structure.
[0063] These elements 22, 23, 24, 25 and 26 are arranged (on-board
wiring) so that a signal supplied from the detector substrate 20A
is supplied to the capacitor 22, resistor 23, analog ASIC 24, ADC
25 and digital ASIC 26 in that order.
[0064] The ASIC substrate 20B includes a connector (spiral contact)
C1 which is connected to the on-board wiring which is connected to
each capacitor 22 to make electrical connections to the detector
substrate 20A and a substrate connector C2 which makes electrical
connections to the data processing apparatus (the unit combination
FPGA which will be described later). Note that the above described
detector substrate 20A also includes the connector C1 connected to
the on-board wiring which is connected to each detector 21.
[0065] (Connection Structure Between Detector Substrate and ASIC
Substrate)
[0066] The connection structure between the detector substrate 20A
and ASIC substrate 20B will be explained.
[0067] The detector substrate 20A and ASIC substrate 20B are
connected not with their respective end faces (ends) facing each
other but by providing an overlap area where both ends overlap with
each other and connecting the connectors C1 in this overlap area as
shown in FIG. 7B. This connection is made in a
detachable/attachable manner using screws for clamping. These
connections are made for the following reason. That is, when the
combined substrate 20 made up of the detector substrate 20A and
ASIC substrate 20B connected (combined) together is supported on
one end (cantilever support) or on both ends in the horizontal
direction, a force which flexes or bends the combined substrate 20
downward is applied to the central area (connection area) of the
combined substrate 20. Here, in the case where both ends are the
connection area where their respective end faces (ends) face each
other, the connection area is easily flexed or bent, which is not
preferable.
[0068] With consideration given to this aspect, this embodiment
connects the detector substrate 20A and ASIC substrate 20B not with
the respective end faces facing each other but by providing the
overlap area so that the areas close to the ends overlap with each
other as described above. This improves toughness against flexure
or bending compared to the connection with the end faces facing
each other, which is preferable. Moreover, improving toughness
against flexure or bending of the combined substrate suppresses
dislocation of the detectors 21 and prevents deterioration of
accuracy of identifying positions at which .gamma.-rays are
generated. As shown in FIG. 2, the camera 11 of the PET apparatus 1
is provided with many detector units 2 (FIG. 10) including the
combined substrate 20 shown in FIGS. 7A-7C in a doughnut shape and
these combined substrates 20 disposed at positions of 3 o'clock and
9 o'clock in the horizontal direction in FIG. 2 are liable to
flexure or bending. Thus, the toughness of the combined substrates
20 against flexure or bending becomes important.
[0069] The detector substrate 20A and ASIC substrate 20B are
electrically connected using the aforementioned overlap area. For
this purpose, a connector C1 (FIG. 7A) which electrically connects
the on-board wiring of both the substrates 20A and 20B is provided
in the respective overlap areas of the detector substrate 20A and
ASIC substrate 20B shown in FIG. 7B. For the connector C1, for
example, a spiral contact (R) is used to improve electrical
connections. The spiral contact (R) is made of a ball-shaped
connection terminal contacting a spiral contactor over a wide area
and provides a characteristic of realizing optimal electrical
connections. Note that when the ball-shaped connection terminal is
provided on the ASIC substrate 20B side, the spiral contactor is
provided on the detector substrate 20A side, and when the
ball-shaped connection terminal is provided on the detector
substrate 20A side, the spiral contactor is provided on the ASIC
substrate 20B side.
[0070] Using such an electrical connection structure between the
detector substrate 20A and ASIC substrate 20B allows signals to be
sent from the detector substrate 20A to the ASIC substrate 20B with
low loss. Note that when loss is small, the energy resolution on
the part of the detectors 21 improves.
[0071] Furthermore, as described above, the detector substrate 20A
and ASIC substrate 20B are connected in a freely
detachable/attachable manner by means of screws, etc. Thus, even
when trouble occurs in the semiconductor radiation detectors 21 or
ASICs 24, 26, all that should be done is just to replace the part
with trouble. Therefore, this eliminates waste that the entire
combined substrate 20 must be replaced due to trouble in that part.
Moreover, since electrical connection between the detector
substrate 20A and ASIC substrate 20B is made by the connector C1
such as the aforementioned spiral contactor (R), connection or
disconnection (combination or dissociation) between the substrates
can be done easily.
[0072] In the above described construction, one detector substrate
20A is connected to the ASIC substrate 20B, but it is also possible
to divide the detector substrate into a plurality of portions. For
example, two detector substrates may be connected to the ASIC
substrate, each consisting of eight horizontal by four vertical
detectors 21. According to this construction, if one detector 21
has trouble, of the two detector substrates, only the one including
the faulty detector 21 needs to be replaced and it is therefore
possible to reduce waste in maintenance (cost reduction).
[0073] (Element Layout)
[0074] Then, the layout of elements such as the detectors 21, ASICs
24, 26 of the combined substrate 20 will be explained with
reference to FIGS. 7A-7C and FIG. 8.
[0075] As shown in FIG. 8, the detector 21 is connected to the
analog ASIC 24 through the connector C1, capacitor 22 and resistor
23 by means of electrical wiring (not shown) and a detection signal
of .gamma.-rays detected by the detector 21 is passed through the
capacitor 22 and resistor 23 by means of the electrical wiring and
processed by the analog ASIC 24. Furthermore, the signal processed
by the analog ASIC 24 is also processed by the ADC 25 and digital
ASIC 26.
[0076] Here, the shorter the length of the circuit and length
(distance) of the wiring, the better, because there is less
influence of noise or less attenuation of a signal. Furthermore,
when simultaneous measurement processing is carried out by the PET
apparatus 1, a shorter circuit or wiring is preferred because its
time delay is smaller (preferable because the accuracy of detection
time is not lost). For this reason, in order of the detector 21,
capacitor 22, resistor 23, analog ASIC 24, ADC 25 and digital ASIC
26 from the center axis of the camera 11 outward in the radius
direction of the camera 11, that is, the elements 21, 22, 23, 24,
25 and 26 are arranged (layout) in this embodiment as shown in FIG.
7A. This order is the same as the signal processing order by the
elements 21, 22, 23, 24, 25 and 26 (see FIG. 8, FIG. 9). That is,
from the center axis of the camera 11 outward, the "detector,
analog integrated circuit, AD converter and digital integrated
circuit" are arranged on the substrate in that order and wired in
the same order. Thus, it is possible to transmit a weak signal
detected by the detector 21 to the analog ASIC 24 with the wiring
length (distance) shortened.
[0077] Since the signal of the analog ASIC 24 is subjected to
processing such as amplification, it is less susceptible to
influences of noise even if the length of wiring from the analog
ASIC 24 onward is long. That is, considering noise, there is no
problem even if the wiring length from the analog ASIC 24 onward is
long. However, with lengthy wiring, there is a delay in signal
transmission and the accuracy of the above described detection time
may deteriorate.
[0078] In this embodiment, since not only the detector 21 but also
the analog ASIC 24 and digital ASIC 26 are included in one combined
substrate 20, the detector 21, analog ASIC 24 and digital ASIC 26
can be arranged in the longitudinal direction of the bed 14, that
is, the direction perpendicular to the body axis of the examinee
subject to an examination, and therefore this eliminates the need
to extend the length of the camera (image pickup apparatus) 11 in
the longitudinal direction of the bed more than necessary. It is
also possible to consider the possibility of arranging the analog
ASIC 24 and digital ASIC 26 outside in the radius direction of the
annularly arranged detector group, and in the longitudinal
direction of the bed 14, but this causes the length of the camera
11 in the longitudinal direction of the bed to become longer than
necessary. Furthermore, semiconductor radiation detectors are used
as the detectors 21, and analog ASIC 24 and digital ASIC 26 are
used as signal processing apparatuses, the length of the combined
substrate 20 in the longitudinal direction is shortened and the
length of the camera 11 in the orthogonal direction can be
shortened significantly compared to the case where a scintillator
is used. Furthermore, the combined substrate 20 is provided with
the detector 21, analog ASIC 24 and digital ASIC 26 in that order
in the longitudinal direction thereof, and therefore it is possible
to shorten the length of the wiring connecting them and simplify
the wiring on the substrate.
[0079] Here, in this embodiment, one analog ASIC 24 is connected to
32 detectors 21 to process signals obtained from the detectors 21.
As shown in FIG. 8 and FIG. 9, one analog ASIC 24 is provided with
32 sets of an analog signal processing circuit (analog signal
processing apparatus) 33 made up of a slow system and fast system.
One analog signal processing circuit 33 is provided for each
detector 21 and connected to one detector 21. Here, the fast system
is provided with a timing pick off circuit 24a to output a timing
signal for identifying a detection time of .gamma.-rays. On the
other hand, the slow system is provided with a polarity amplifier
(linear amplifier) 24c, a band pass filter (waveform shaping
apparatus) 24d and a peak hold circuit (peak value holding
apparatus) 24e connected in this order for the purpose of
calculating a peak value of the detected .gamma.-rays. Note that
the slow system is named "slow" because it takes a certain degree
of processing time to calculate a peak value. Reference numeral 24b
denotes a charge amplifier (preamplifier). A .gamma.-ray detection
signal output from the detector 21 and passed through the capacitor
22 and resistor 23 is amplified at the charge amplifier 24b and
polarity amplifier 24c. The amplified gamma-ray detection signal is
passed through the band pass filter 24d and input to the peak hold
circuit 24e. The peak hold circuit 24e holds a maximum value of the
detection signal, that is, the peak value of a .gamma.-ray
detection signal proportional to energy of the detected
.gamma.-rays. One analog ASIC 24 is an LSI which integrates 32 sets
of analog signal processing circuits 33.
[0080] The capacitor 22 and resistor 23 can also be provided inside
the analog ASIC 24, but this embodiment arranges the capacitor 22
and resistor 23 outside the analog ASIC 24 for reasons such as
obtaining an appropriate capacitance and appropriate resistance and
reducing the size of the analog ASIC 24. Note that the capacitor 22
and resistor 23 are preferably disposed outside because in this way
variations in the individual capacitance and resistance are
reduced.
[0081] In the analog ASIC 24 shown in FIG. 8, the output of the
slow system of this analog ASIC 24 in this embodiment is designed
to be supplied to an ADC (analog/digital converter) 25. Moreover,
the output of the fast system of the analog ASIC 24 is designed to
be supplied to the digital ASIC 26.
[0082] The analog ASIC 24 and each ADC 25 are connected via one
wire which sends slow system signals corresponding to 8 channels
all together. Furthermore, each analog ASIC 24 and digital ASIC 26
are connected via 32 wires which send 32-channel fast system
signals one by one. That is, one digital ASIC 26 is connected to
four analog ASICs 24 via a total of 128 wires.
[0083] The output signal of the slow system output from the analog
ASIC 24 is an analog peak value (maximum value of the graph shown
in FIG. 4). Furthermore, the output signal of the fast system
output from the analog ASIC 24 to the digital ASIC is a timing
signal showing the timing corresponding to the detection time. Of
these signals, the peak value which is the slow system output is
input to the ADC 25 via the wire (wire uniting 8 channels into one)
connecting the analog ASIC 24 and ADC 25 and converted to a digital
signal by the ADC 25. The ADC 25 converts a peak value to, for
example, an 8-bit (0 to 255) digital peak value (e.g., 511
KeV.fwdarw.255). On the other hand, a timing signal which is a slow
system output is supplied to the digital ASIC 26 via the above
described wire connecting the analog ASIC 24 and digital ASIC
26.
[0084] The ADC 25 sends the digitized 8-bit peak value information
to the digital ASIC 26. For this purpose, each ADC 25 and digital
ASIC 26 are connected via a wire. For example, since there are
sixteen ADCs 25 on both sides, the digital ASIC 26 is connected to
the ADC 25 via a total of sixteen wires. One ADC 25 processes
signals corresponding to 8 channels (signals corresponding to eight
detection elements). The ADC 25 is connected to the digital ASIC 26
via one wire for transmission of an ADC control signal and one wire
for transmission of peak value information.
[0085] As shown in FIG. 9, the digital ASIC 26 comprises a
plurality of packet data generation apparatuses 34 including eight
time decision circuits (time information generation apparatuses) 35
and one ADC control circuit (ADC control apparatus) 36, and a data
transfer circuit (data transmission apparatus) 37, and integrates
all these elements into one LSI. All the digital ASICs 26 provided
for the PET apparatus 1 receive a 500 MHz clock signal from a clock
generation apparatus (crystal oscillator) (not shown) and operates
synchronously. The clock signal input to each digital ASIC 26 is
input to the respective time decision circuits 35 in all the packet
data generation apparatuses 34. One time decision circuit 35 is
provided for each detector 21 and receives a timing signal from the
timing pick off circuit 24a of the corresponding analog signal
processing circuit 33. The time decision circuit 35 determines the
detection time of .gamma.-rays based on the clock signal when the
timing signal is input. Since the timing signal is based on the
fast system signal of the analog ASIC 24, a time close to a real
detection time can be set as the detection time (time information).
The ADC control circuit 36 receives a timing signal for the timing
at which .gamma.-rays are detected from the time decision circuit
35 and identifies the detector ID. That is, the ADC control circuit
36 stores a detector ID corresponding to each time decision circuit
35 connected to the ADC control circuit 36 and can identify, when
time information is input from a certain time decision circuit 35,
the detector ID corresponding to the time decision circuit 35. This
is possible because one time decision circuit 35 is provided for
each detector 21. Moreover, after inputting the time information,
the ADC control circuit 36 outputs an ADC control signal including
detector ID information to the ADC 25. The ADC 25 outputs the peak
value information output from the peak hold circuit 24e of the
analog signal processing circuit 33 corresponding to the detector
ID by converting it to a digital signal. This peak value
information is input to the ADC control circuit 36. The ADC control
circuit 36 adds the peak value information to the time information
and detector ID to create packet data. The ADC control circuit 36
has the functions of the ADC control apparatus for controlling the
ADC 25 and the information combination apparatus for combining the
detector ID information (detector position information), time
information and peak value information. The information combination
apparatus outputs combination information (packet information)
which is digital information including those three kinds of
information. The packet data (including detector ID, time
information and peak value information) output from the ADC control
circuit 36 of each packet data generation apparatus 34 is input to
the data transfer circuit 38.
[0086] The data transfer circuit 38 sends packet data which is
digital information output from the ADC control circuit 36 of each
packet data generation apparatus 34 to the integrated circuit (unit
combination FPGA (Field Programmable Gate array)) 31 for unit
combination provided for the housing 30 of the detector unit 2
(FIG. 10, FIG. 11) which houses twelve combined substrates 20, for
example, periodically. The unit combination FPGA (hereinafter
referred to as "FPGA") 31 sends the digital information to the data
processing apparatus 12 through an information transmission wire
connected to the connector 38.
[0087] Since the ADC 25 converts the peak value information output
from the peak hold circuit 24e corresponding to the detector ID
information included in a control signal output from the ADC
control circuit 36 to a digital signal, one ADC 25 is provided for
a plurality of analog signal processing circuits 33 in one analog
ASIC 24. Therefore, there is no need to provide one ADC 25 for each
of a plurality of analog signal processing circuits 33 and it is
possible to thereby significantly simplify the circuit construction
of the ASIC substrate 20B. Also one information combination
apparatus for generating combination information is enough for a
plurality of analog signal processing circuits 33 in one analog
ASIC 24, which can simplify the circuit construction of the digital
ASIC 26. Moreover, only one ADC control apparatus for identifying
detector IDs needs to be provided for a plurality of analog signal
processing circuits 33 in one analog ASIC 24, simplifying the
circuit construction of the digital ASIC 26.
[0088] In this way, packet data output from the digital ASIC 26 and
including detector IDs for uniquely identifying (1) peak value
information, (2) determined time information and (3) detector 21,
one by one is sent to the next data processing apparatus 12 (see
FIG. 1) through an information transmission wire. The simultaneous
measuring apparatus 12A of the data processing apparatus 12 carries
out simultaneous measuring processing (when two .gamma.-rays with
predetermined energy are detected with a time window with a set
time, this processing regards these .gamma.-rays as a pair of
.gamma.-rays generated by annihilation of one positron) based on
the packet data sent from the digital ASIC 26, counts the
simultaneously measured pair of .gamma.-rays as one .gamma.-ray and
identifies the positions of the two detectors 21 which have
detected the pair of .gamma.-rays using those detector IDs. When
there are three or more .gamma.-rays detection signals detected
within the above described time window (when there are three or
more detected detectors 21 which have detected .gamma.-rays) the
data processing apparatus 12 identifies the two detectors 21 into
which .gamma.-rays are introduced first out of three or more
detectors 21 using peak value information, etc., on those
.gamma.-ray detection signals. The identified one pair of detectors
21 are simultaneously measured and one count value is generated.
Furthermore, the tomographic information creation apparatus 12B of
the data processing apparatus 12 creates tomographic information on
the examinee at the position where radiopharmaceuticals are
concentrated, that is, position of malignant tumor, using count
values obtained by simultaneous measurement and position
information on the detectors 21. This tomographic information is
displayed on the display apparatus 13. Information such as the
above described digital information, count value obtained by
simultaneous measurement and position information on the detectors
21 and tomographic information are stored in the storage apparatus
of the data processing apparatus 12.
[0089] According to the above described explanations, the detector
substrate 20A includes the detectors 21 and the ASIC substrate 20B
includes the capacitor 22, resistor 23, analog ASIC 24, ADC 25 and
digital ASIC 26. However, the detector substrate (first substrate)
20A may include the detector 21, capacitor 22, resistor 23 and
analog ASIC 24, etc., and the ASIC substrate (second substrate) 20B
may include the ADC 25 and digital ASIC 26, etc. By the detector
substrate 20A including the detectors 21 and analog ASIC 24, the
distance (wire length) between the detector 21 and analog ASIC 24
can be further shortened. Thus, it is possible to further reduce
influences of noise.
[0090] Furthermore, the combined substrate 20 may include three
substrates (detector substrate 20A, analog ASIC substrate and
digital ASIC substrate) and they may be connected in a
detachable/attachable manner through their respective
connectors.
[0091] In this case, of the three substrates, the detector
substrate 20A includes the detectors 21, the analog ASIC substrate
includes the capacitor 22, resistor 23 and analog ASIC 24 and the
digital ASIC substrate includes the ADC 25 and digital ASIC 26.
This structure separates the substrate incorporating the analog
circuit from the substrate incorporating the digital circuit to
thereby prevent noise on the digital circuit side from entering the
analog circuit. Furthermore, this structure separates the substrate
incorporating the analog ASIC from the substrate incorporating the
digital ASIC and connects the two substrates using a
detachable/attachable connector, and therefore even when only the
digital ASIC malfunctions, only the digital ASIC substrate needs to
be replaced. In this way, this structure can further reduce
waste.
[0092] In the above explanations, the substrate body 20a (detector
substrate 20A) for mounting the detectors 21 is different from the
substrate body 20b (ASIC substrate 20B) for mounting the ASICs 24,
26. Thus, when, for example, both ASICs are soldered to a substrate
by means of a BGA (Ball Grid Array) using reflow, only the ASIC
substrate can be soldered and this is preferable because the
semiconductor radiation detector 21 need not be exposed to a high
temperature. Of course, it is also possible to arrange all the
elements 21 to 26 on the same substrate and use no connector
C1.
[0093] <<Detector Unit; Unit Construction Through Housing of
Combined Substrate>>
[0094] Next, a unit construction by housing the above described
combined substrate 20 in the housing 30 will be explained. This
embodiment constructs a detector unit (twelve substrate units) 2 by
housing twelve combined substrates 20 in the housing (frame) 30.
The camera 11 of the PET apparatus 1 has a structure in which 60 to
70 detector units 2 are arranged in the circumferential direction
in a detachable/attachable manner (see FIG. 12B) so as to
facilitate maintenance and examination (see FIG. 2).
[0095] (Housing in Housing)
[0096] As shown in FIG. 10, the detector unit 2 is provided with a
housing 30, etc., for housing or holding the above described 12
combined substrates 20, a high-voltage power supply PS for
supplying a charge collecting voltage to these 12 combined
substrates 20, the above described FPGA 31, signal connectors for
exchanging signals with the outside and power connectors for
receiving a power supply from the outside.
[0097] As shown in FIG. 10 and FIG. 11, the combined substrates 20
are housed in the housing 30, arranged in three rows in the depth
direction (longitudinal direction of the bed 14) without
overlapping with one another and in four rows in the width
direction (circumferential direction of the camera 11). Namely, one
housing 30 houses twelve combined substrates 20. To realize such
housing, a guide member 39 consisting of four rows of guide grooves
(guide rails) G1 extending in the depth direction and arranged at
appropriate intervals in the circumferential direction is disposed
in the housing 30 and fitted at the upper end of the housing
(cover) 30. The guide member 39 has an opening 40 opposed to each
connector C3 of a ceiling plate 30a in the portion of each guide
groove G1. Furthermore, a bottom plate 30b of the housing 30 is
provided with four guide members 41 having one guide groove (guide
rail) G2 extending in the depth direction arranged at appropriate
intervals in the circumferential direction (see FIG. 11). The guide
grooves G1, G2 have a depth corresponding to a capacity of housing
three combined substrates 20. An end of the combined substrate 20
on the ASIC substrate 20B side is housed in the guide groove G1 and
an end of the combined substrate 20 on the detector substrate 20A
side is housed in the guide groove G2. Three combined substrates 20
are held in the depth direction of the guide grooves G1, G2. Note
that since the end of the combined substrate 20 on the ASIC
substrate 20B side and the other end on the detector substrate 20A
side are designed to be slidable in the guide grooves G1, G2, it is
possible to easily position the combined substrates 20 at
predetermined locations by sliding them in the guide grooves G1, G2
with fingers, for example. In this case, each substrate connector
C2 is positioned in the portion of each opening 40. After a
predetermined number of combined substrates 20 are arranged in the
housing 30, the ceiling plate 30a is attached at the top end of the
housing 30 in a detachable/attachable manner using screws, etc.
Each connector C3 fitted in the ceiling plate 30a is inserted in
the corresponding opening 40 and connected to the corresponding
substrate connector C2. The terms "upper" and "lower" sections of
the housing 30 are applicable when the housing 30 is removed from
the camera 11, and when the housing 30 is mounted in the camera 11
as shown in FIG. 2, the upper and lower sections may be inverted or
turned 90 degrees to be "right" and "left" sections or located
diagonally.
[0098] As shown in FIG. 11, the ceiling plate 30a of the housing 30
is provided with not only the four rows of guide grooves G1 but
also FPGA 31 and connector 38. The connector 38 is connected to the
FPGA 31. The FPGA 31 is programmable in the field. In this aspect,
the FPGA 31 is different from the ASIC in that it is not
programmable. Therefore, as with this embodiment, even if the
number or type of the combined substrates 20 to be housed changes,
the FPGA 31 can be programmed in the field to be adaptable to
changes in the number of substrates appropriately.
[0099] Since the detectors 21 using CdTe as the semiconductor
material S in this embodiment generate charge in reaction with
light, the housing 30 is made of a material having light shielding
properties such as aluminum and an alloy of aluminum and designed
in such a way as to eliminate gaps through which light enter. That
is, the housing 30 is constructed to secure light shielding
properties. If, for example, light shielding properties are secured
by other means, the housing 30 itself need not be provided with
light shielding properties and the housing 30 can be a frame
(framework) to hold the detectors 21 in a detachable/attachable
manner (e.g., no light shielding plane member (panel), etc., is
required).
[0100] As shown in FIG. 12A, the detector unit 2 is mounted via a
unit support member 2A. Furthermore, as shown in FIG. 12B, the
detector unit 2 is mounted in the camera 11 with one end supported
by the unit support member 2A. The unit support member 2A has a
hollow disk (doughnut) shape and is provided with many windows (as
many as the detector units 2 to be mounted) in the circumferential
direction of the camera 11. In order to support the detector units
2 at one end, a flange portion serving as a stopper is provided on
the front side in the axial direction of the body of the housing 30
of the detector unit 2. Note that the flange portions inside in the
circumferential direction become obtrusive when the detector units
2 are arranged as dense as possible in the circumferential
direction. Therefore, it is possible to remove the obtrusive flange
portions from the housing 30 and allow the flange portions outside
in the circumferential direction to remain. Or it is also possible
to provide another unit support member 2A and support both ends of
the detector unit 2 by the two unit support members 2A.
[0101] In order to mount the detector units 2 in the unit support
member 2A, this embodiment allows many detectors 21 to be mounted
in the camera 11 at a time. This can considerably shorten the time
of mounting the detectors 21 in the camera 11. Furthermore, packet
data (all packet data for all detectors 21 of a combined substrate
20) output from the data transfer apparatus 38 of all the combined
substrates 20 in the detection unit 2 is sent from the unit
combination FPGA 31 provided in the detection unit 2 to the data
processing apparatus 12. In this way, the number of wires through
which packet data is transmitted to the data processing apparatus
12 in this embodiment is also significantly reduced compared to the
case where packet data is sent from each data transfer apparatus 38
of the combined substrate 20 to the data processing apparatus
12.
[0102] When the detector units 2 is mounted in the camera 11, a
cover 11a is removed to make the unit support member 2A exposed so
that the detector units 2 are inserted until the detector units 2
touch the flange portions. When the detector units 2 are inserted
and fitted, connections between the camera 11 and the detector
units 2 are made, and signals and power supply are connected
between the camera 11 and the detector units 2.
[0103] (Power Supply)
[0104] Then, the high-voltage power supply apparatus PS for
supplying a charge collection voltage will be explained. As shown
in FIG. 10, the detector unit 2 provides the high-voltage power
supply apparatus PS for supplying a charge collection voltage to
each detector 21 in a space formed inside the housing 30 on the
back of the FPGA 31. This high-voltage power supply apparatus PS
receives a low voltage power supply, boosts the voltage to 300 V
using a DC-DC converter (means for boosting the voltage, which is
not shown) and supplies the voltage to each detector 21. 64
detectors 21 are provided per one combined substrate 20 (=detector
substrate 20A) on one side, and 128 on both sides. Twelve such
combined substrates 20 are housed in one housing 30 (that is, one
detector unit 2). Thus, the high-voltage power supply apparatus PS
supplies voltages to 128.times.12=1536 detectors 21.
[0105] Conventionally, a supply voltage of 300 V with extremely
small fluctuations is supplied from a precision power supply
apparatus in a remote place, but since (1) when the distance from
the precision power supply apparatus increases, a wider insulating
structure for high voltage wiring is required (the insulating
distance increases) and (2) the voltage fluctuates due to a
temperature variation of the detectors 21, there is a problem that
supplying a precise voltage from the precision power supply
apparatus does not necessarily result in a precise voltage in the
part of the target detectors 21.
[0106] Furthermore, to facilitate maintenance and examination, it
is also possible to consider providing the detector unit 2
according to this embodiment with a power connector (not shown) and
removing a high-voltage power line extending from the precision
power supply apparatus at this power connector. That is, according
to this embodiment, it is possible to consider supplying a
high-voltage power supply to the detector units 2 from outside the
units 2 via power connectors. However, in the case of a high
voltage of 300 V, this results in a problem that the size of the
power connector increases in addition to the above described
problem of insulation.
[0107] According to this embodiment, the high-voltage power supply
apparatus PS built in the detector unit 2 is connected to an
external low voltage (5 to 15 V) DC power supply through the power
connector 42 and connector 38 provided on the ceiling plate 30a via
power wiring. A high-voltage terminal of the high-voltage power
supply apparatus PS is connected to twelve connectors C3 provided
on the ceiling plate 30a through the connector 43 provided on the
ceiling plate 30a and connected to electrodes C of the respective
detectors 21 provided on the substrate body 20a through the
connector C2 of the respective combined substrates 20, power wiring
(not shown) in the substrate body 20b, connector C1 and power
wiring (not shown) in the substrate body 20a. The connectors C1, C2
include not only connectors for transmitting output signals of the
detectors 21 but also connectors for power wiring. Since the
high-voltage power supply apparatus PS boosts a low voltage applied
from the power supply to 300 V using a DC-DC converter, it is
possible to reduce the high-voltage section and thereby shorten the
insulation distance. That is, this eliminates the necessity for
using high-voltage wiring for a portion from the connector 42 to
the DC power supply. It also facilitates maintenance, etc. For the
problem with voltage fluctuations, this embodiment provides not the
high-precision power supply apparatus but the high-voltage power
supply apparatus PS having accuracy according to a temperature
fluctuation of the voltage. This eliminates the necessity for a
high-precision power supply. Furthermore, since it is a low voltage
that is received from an external power supply, it is possible to
use a small power connector to be provided for the connector 38.
Using the small power connector increases the degree of freedom in
the layout. Furthermore, since the high-voltage power supply
apparatus PS is arranged in a space formed in the housing 30 on the
back side of the FPGA 31, the arrangement of the high-voltage power
supply apparatus PS in the housing 30 makes the detector unit 2
more compact instead of upsizing. It is also possible to directly
connect the high-voltage power supply apparatus PS to the power
wiring provided on the substrate body 20a through the connector,
without the ceiling plate 30a. Furthermore, the power connector can
also be separated from the output signal connector of the detector
21. This prevents noise from entering the signal wiring from the
power supply system.
[0108] Furthermore, by reducing a supply voltage to the detector
unit 2, it is possible to supply power to the high-voltage power
supply apparatus PS at a low voltage through the unit combination
FPGA 31 as with power supplies to the ASICs 24, 26.
[0109] Furthermore, supplying power using the high-voltage power
supply apparatus PS eliminates the necessity for insulation from
the housing (GND).
[0110] The voltage supplied from the FPGA 31 to the high-voltage
power supply apparatus PS is boosted to 300 V by a DC-DC converter
(not shown) in the high-voltage power supply apparatus PS and after
boosting, passed through the ceiling plate 30a of the housing 30
and supplied from ASIC substrate 20B.fwdarw.detector substrate
20A.fwdarw.each detector 21 for each combined substrate 20. That
is, the housing 30 (ceiling plate 30a) is provided with wiring for
voltage supply (not shown) for supplying a voltage from the
high-voltage power supply apparatus PS to each combined substrate
20. Furthermore, each combined substrate 20 is provided with wiring
for voltage supply which supplies a voltage supplied from the
high-voltage power supply apparatus PS to each detector 21 via the
substrate connector C2.
Embodiment 2
[0111] A nuclear medicine diagnostic apparatus according to another
embodiment will be explained. The nuclear medicine diagnostic
apparatus of this embodiment is single photon emission computer
tomography (SPECT) apparatus.
[0112] This SPECT apparatus 51 will be explained using FIGS. 13 to
15. The SPECT apparatus 51 is provided with a pair of radiation
detection blocks 52, a rotary holder (body of rotation) 57, a data
processing apparatus 12A and a display apparatus 13. The radiation
detection blocks 52 are disposed at two positions with a
180.degree. difference in the circumferential direction of the
rotary holder 57. More specifically, the respective unit support
members 56 of the radiation detection blocks 52 are mounted on the
rotary holder 57 with a 180.degree. difference in the
circumferential direction. A plurality of detector units 2A each
including twelve combined substrates 53 are mounted on the
respective unit support members 56 in a detachable/attachable
manner. Thus, the detectors 21 are supported by the unit support
member. The construction of each detector unit 2A is the same as
that of the detector unit 2 according to Embodiment 1 except the
construction of the combined substrate 53.
[0113] The combined substrate 53 includes a detector substrate 20A
and an ASIC substrate 53B as with the above described combined
substrate 20 (FIG. 14). The detectors 21 at one end of each
detector substrate 20A are arranged facing the bed 14. A collimator
55 made of a radiation shielding member (e.g., lead, tungsten,
etc.) is provided on each radiation detection block 52. Each
collimator 55 forms many radiation passages through which radiation
(e.g., .gamma.-rays) passes. These radiation passages are provided
in a one-to-one correspondence with the detectors 21 positioned at
one end of all the detector substrates 20A of one radiation
detection block 52. All the combined substrates 53 and collimators
55 are arranged within a light/electromagnetic shield 54 mounted on
the rotary holder 57. The collimator 55 is mounted in the
light/electromagnetic shield 54. The light/electromagnetic shield
54 cuts off influences of electromagnetic waves other than
.gamma.-rays on the detectors 21, etc.
[0114] When the bed 14 on which an examinee administered with
radiopharmaceuticals is laid is moved, the examinee is moved
between the pair of radiation detection blocks 52. When the rotary
holder 57 is rotated, the detector units 2A of each radiation
detection block 52 revolve around the examinee. .gamma.-rays
emitted form an area in the body of the examinee where
radiopharmaceuticals are concentrated (e.g., affected area) C pass
through the radiation passages of the collimator 55 and are
introduced into the corresponding detectors 21. The detectors 21
output .gamma.-rays detection signals. These .gamma.-ray detection
signals are processed by analog ASIC 24A and digital ASIC 26A,
which will be described later.
[0115] The construction of the detector substrate 20A used in this
embodiment (Embodiment 2) is the same as that in Embodiment 1 and
therefore the explanations will be omitted in this embodiment. The
ASIC substrate 53B making up the combined substrate 53 will be
explained using FIGS. 14 and 15. As with the combined substrate 20,
the ASIC substrate 53B connected to the detector substrate 20A
through the connector C1 includes a capacitor 22 and a resistor 23,
four analog ASICs 24A and one digital ASIC 26A for each detector
21.
[0116] One analog ASIC 24A is provided with 32 sets of analog
signal processing circuits (analog signal processing apparatuses)
33A having a slow system and fast system. One analog signal
processing circuit 33A is provided for each detector 21. Here, the
fast system includes a trigger output circuit 24f which outputs a
trigger signal for specifying detection of .gamma.-rays. As with
the analog ASIC 24, the slow system is provided with a charge
amplifier 24b, a polarity amplifier 24c, a band pass filter 24d and
a peak hold circuit 24e connected in this order. One analog ASIC
24A integrates 32 sets of analog signal processing circuits 33A
into one LSI. A .gamma.-ray detection signal which is output from
the detector 21 and has passed through the capacitor 22 and
resistor 23 are guided through the charge amplifier 24b, polarity
amplifier 24c and band pass filter 24d and input to the peak hold
circuit 24e. The peak hold circuit 24e holds a peak value of the
.gamma.-ray detection signal. The .gamma.-ray detection signal
output from the band pass filter 24d is input to the trigger output
circuit 24f. The trigger output circuit 24f outputs a trigger
signal when a .gamma.-ray detection signal at a set level or higher
is input to remove influences of noise.
[0117] The digital ASIC 26A includes a packet data generation
apparatus 34A and a data transfer circuit 37 and integrates them
into one LSI. The above described trigger signal is input to the
ADC control circuit 36A of the packet data generation apparatus
34A. All the digital ASICs 26A provided on the SPECT apparatus 51
receive a 64 MHz clock signal from a clock generation apparatus
(crystal oscillator) (not shown) and operate synchronously. The
clock signal input to each digital ASIC 26A is input to the
respective ADC control circuits 36A in all the packet data
generation apparatuses 34A. The ADC control circuit 36A identifies
the detector ID when the trigger signal is input. That is, the ADC
control circuit 36A stores a detector ID for each trigger output
circuit 24f connected to the ADC control circuit 36A and can
identify, when a trigger signal is input from a certain trigger
output circuit 24f, the detector ID corresponding to the trigger
output circuit 24f. The ADC control circuit 36A outputs an ADC
control signal including the detector ID information to the ADC 25.
The ADC 25 converts the peak value information output from the peak
hold circuit 24e of the analog signal processing circuit 33A
corresponding to the detector ID to a digital signal and outputs
it. This peak value information is input to the ADC control circuit
36. The ADC control circuit 36A adds the peak value information to
the detector ID to generate packet data. The packet data (including
detector ID and peak value information) which is the digital
information output from the ADC control circuit 36A of each packet
data generation apparatus 34A is input to the data transfer circuit
37. The data transfer circuit 37 sends the packet data output from
each ADC control circuit 36A to the unit combination FPGA 31 of the
detector unit 2A periodically. The unit combination FPGA 31 outputs
the digital information to the information transmission wiring
connected to the connector 38.
[0118] Packet data output from the unit combination FPGA 31 is sent
to the data processing apparatus 12A. A rotation angle detected by
an angle gauge (not shown) connected to the rotation shaft of a
motor (not shown) for rotating the rotary holder 57 is input to the
data processing apparatus 12A. This rotation angle indicates the
rotation angle of each radiation detection block 52 and more
specifically indicates the rotation angle of each detector 21.
Based on this rotation angle, the data processing apparatus 12A
determines the position (position coordinates) of each revolving
detector 21 on the revolving orbit. In this way, the position
(position coordinates) of the detector 21 when .gamma.-rays are
detected is calculated. Based on the calculated position of the
detector 21, the data processing apparatus 12A counts .gamma.-rays
whose peak value information reaches and exceeds a set value. This
counting is performed on each area obtained by dividing the
revolving circle into 0.5.degree. portions relative to the
rotational center of the rotary holder 57. The peak value
information is an accumulated value of peak values of respective
.gamma.-ray detection signals of a plurality of detectors 21 (four
detectors 21 arranged on a straight line in FIG. 7A) positioned on
an extension of the radiation passage of the collimator 55. Using
the position information of the detectors 21 and count value (count
information) of .gamma.-rays when .gamma.-rays are detected, the
data processing apparatus 12A creates tomographic information on a
position at which radiopharmaceuticals are concentrated, that is,
position of malignant tumor of the examinee. This tomographic
information is displayed on the display apparatus 13. Information
such as the above described packet information, count value
obtained by simultaneous measurement, position information of the
detector 21 and tomographic information are stored in the storage
apparatus of the data processing apparatus 12.
[0119] The foregoing embodiments have described the PET apparatus 1
and SPECT apparatus 51, but the present invention is also
applicable to a .gamma. camera. Functional images obtained from the
.gamma. camera are two-dimensional and the .gamma. camera is
provided with a collimator for regulating angles of incidence of
.gamma.-rays. Moreover, it is also possible to adopt a construction
of a nuclear medicine diagnostic apparatus combining the PET
apparatus 1 and SPECT apparatus 51, and an X-ray CT.
[0120] Mounting (housing) of the detector unit 2 in the camera 11
is not limited to the mounting using the above described unit
support member 2A, but any mounting/housing means or method can be
used.
[0121] It should be further understood by those skilled in the art
that although the foregoing description has been made on
embodiments of the invention, the invention is not limited thereto
and various changes and modifications may be made without departing
from the spirit of the invention and the scope of the appended
claims.
* * * * *