JP2008286560A - Crystal element assembly, electric circuit therefor, nuclear medicine diagnosis apparatus using them, and energization control method - Google Patents

Abstract

An object of the present invention is to provide a nuclear medicine diagnostic apparatus and a nuclear medicine diagnostic method excellent in energy resolution.
A nuclear medicine diagnostic apparatus is formed by bonding a semiconductor element S and metal conductive members 22A, 22B, 23A, and 23B with a conductive adhesive 24 composed of conductive particles and a resin binder. A structure having a structure in which charges generated when γ rays are incident on the semiconductor element S are extracted as signals from the conductive members 22A, 22B, 23A, and 23B through the detection circuit 40 via the conductive adhesive 24; An energization control means 30 is provided for causing a current larger than the current due to the charge generated when the line is incident to flow through at least the conductive adhesive 24.
[Selection] Figure 5

Description

  The present invention relates to a crystal element assembly in which a conductive member for connecting to an external electric circuit in order to use the crystal element as a functional element, and the electrical connection between the conductive member and the crystal element in the crystal element assembly. ELECTRIC CIRCUIT FOR CRYSTAL ELEMENT ASSEMBLY FOR MAINTAINING GOOD CONDUCTIVITY, NUCLEAR MEDICAL DIAGNOSIS DEVICE USING THE CRYSTAL ELEMENT ASSEMBLY AS RADIATION DETECTOR, AND METHOD FOR CONTROLLING ELECTRIC CIRCUIT FOR ELECTRIC CIRCUIT FOR CRYSTAL ELEMENT ASSEMBLY Is.

In recent years, gamma cameras, single photon emission tomography devices (SPECT (Single Photon Emission Computed Tomography) devices), and positron emission tomography devices (PET (Positron Emission Tomography)) have been used as devices that apply radiation measurement technology to the medical field. ) Devices) are known, and most of the radiation detectors used in these devices are a combination of a scintillator and a photomultiplier tube. For example, Patent Document 1 discloses that a semiconductor is used instead of a scintillator.
In such a semiconductor radiation detector, electrons and holes are generated by the photoelectric effect caused by the γ rays incident on the crystal of the semiconductor element, and moved by an electric field generated by an externally applied voltage, and the electric charge is transferred to the outside. Since the amount of charge is proportional to the energy of radiation, the energy of radiation can be accurately known by accurately measuring the amount of charge. Therefore, the semiconductor has an advantage that the radiation energy can be measured more accurately than the combination of the scintillator and the photomultiplier tube.

  By the way, as a semiconductor for detecting radiation, particularly γ-rays, the effective atomic number is large in terms of sensitively measuring γ-rays used in nuclear medicine diagnostic apparatuses such as gamma cameras, SPECT apparatuses, and PET apparatuses. Methods using cadmium telluride (CdTe) or zinc cadmium telluride (CdZnTe), which are operable at room temperature in terms of handling, are known. Here, silicon and germanium are also used, but these are often used for X-ray detection. As radioactive materials used for nuclear medicine diagnosis, for example, technetium and fluorine 18 are known, and these have high transmission power because their energy is 141 keV or 511 keV. For this reason, it is advantageous to use a substance having a large atomic number for the radiation detector to absorb a large amount of γ-rays as a signal. In this respect, CdTe and CdZnTe have an advantage that they have an excellent ability to absorb γ rays because they have a large average atomic number.

  When a semiconductor crystal such as CdTe or CdZnTe is used as a radiation detection element of a radiation detector, an anode electrode and a cathode electrode are provided on the semiconductor crystal. A conductive thin film is usually used for the anode electrode and the cathode electrode, and a conductive member is connected to the anode electrode and the cathode electrode in order to connect the anode electrode and the cathode electrode to an external circuit. At this time, since CdTe and CdZnTe are brittle in crystals and insufficient in heat resistance, bonding is required when connecting an anode electrode and a cathode electrode formed by vapor deposition or sputtering to a CdTe or CdZnTe semiconductor crystal and a conductive member. A conductive adhesive is used as the member.

The conductive adhesive is a mixture of conductive flakes (conductive particles) such as silver called a filler in a resin binder such as epoxy, and the fillers come into contact with each other when the resin shrinks during the resin curing process. Thus, conductivity is expressed. Therefore, the joining by the conductive adhesive is different from, for example, the connection relationship between solder and copper, and is rather close to a point contact.
On the other hand, a radiation detection element using a semiconductor crystal such as CdTe or CdZnTe is usually applied with a bias voltage of about several tens to several hundreds volts at the time of radiation measurement. At that time, a dark current around nanoampere is applied to the radiation detection element. For example, when γ rays are incident, a pulse current signal of several tens to several hundreds of nanoamperes is superimposed on the dark current and flows for a moment.
JP 2004-317140 A

Actually, a conductive thin film of an anode electrode and a cathode electrode is formed on the opposing surfaces of the conductive member and the CdTe semiconductor crystal, and the anode electrode and the cathode electrode are faced to each other as a unit. A conductive member was interposed between them, and the anode and cathode electrodes and the conductive member were bonded together with a conductive adhesive, and the units were laminated to produce a semiconductor radiation detector. However, when this manufactured semiconductor radiation detector was experimentally used for several months, in some semiconductor radiation detectors, the noise of the detection signal increased and the energy resolution decreased.
This is because conduction becomes unstable due to a decrease in the shrinkage of the conductive adhesive that bonds the conductive member, the anode electrode, and the cathode electrode after long-term use, or oxidation of the conductive material of the filler. It was inferred to be caused.

  On the other hand, the charge generated in the semiconductor radiation detector due to the input of radiation is a small amount of several femtocoulombs to several tens of femtocoulombs, and the current value is less than microamperes. For this reason, when such a weak signal passes through the unstable part of the conductivity, there is a possibility that the signal may be distorted, which appears as deterioration of the semiconductor radiation detector, leading to a decrease in energy resolution. There was also.

  Accordingly, an object of the present invention is to provide a crystal element assembly having good electrical continuity at the time of measurement, an electric circuit therefor, and a nuclear medicine diagnostic apparatus excellent in stable energy resolution using the crystal element assembly as a radiation detector. And a method of controlling energization of an electric circuit for a crystal element assembly.

  In order to achieve the above-described object, the crystal element assembly according to the first invention of the present application is external to each of the first and second conductive thin films formed on the surface of the crystal element which is a functional element. At least two conductive members for electrical connection to the first and second conductive thin films by means of a conductive adhesive composed of conductive particles and a binder. And an adhesive structure formed by bonding the first or second conductive thin film.

  A second invention of the present application uses the crystal element assembly of the first invention, and a conductive member having a current path between two adhesive structures on one conductive thin film of the first and second conductive thin films The first energization state of the first electroconductive thin film, the crystal element, and the energization control means capable of selecting the second energization state of the conductive member in which the second electroconductive thin film becomes a current path are combined. It is characterized by that.

  According to a third invention of the present application, the crystal element assembly of the first invention is used as a radiation detector, and at the time of radiation measurement, the energization control means is made to select the first energization state so that the crystal element acts as a radiation detection element. The electric charge generated at this time is taken out as an electrical signal from the conductive member connected to the first and second conductive thin films, and when the radiation is not measured, the second energized state is selected to act as a radiation detection element. It is characterized in that a current larger than the current of the electric signal due to the electric charge generated at the time is caused to flow between the adhesive structures on one conductive thin film.

According to the first to third aspects of the present invention, the conductive particles of the conductive adhesive material in the adhesive structure portion are re-coupled or conductive by the large current that flows when the energization control unit selects the second energization state. The oxide film formed on the adhesive can be broken, and the conductivity of the conductive adhesive can be stabilized.
In addition, according to the third aspect of the present invention, the energization in the second energization state by the energization control means is performed before the voltage is applied to the radiation detection element for radiation measurement, so that the occurrence of failure due to unstable conduction is effectively achieved. The radiation detection performance can be maintained over a long period of time.
In addition, this invention includes the electricity supply control method of the electric circuit for crystal element assemblies.

  According to the present invention, a crystal element assembly having good electrical continuity during measurement, an electric circuit therefor, a nuclear medicine diagnostic apparatus excellent in stable energy resolution using this crystal element assembly as a radiation detector, and An electrical circuit energization control method for a crystal element assembly can be provided.

Next, a PET apparatus according to a preferred embodiment of the present invention will be described in detail with reference to the drawings as an example.
<< First Example of Semiconductor Radiation Detector >>
First, a detailed configuration of the semiconductor radiation detector used in the present embodiment will be described.
FIG. 1A is a perspective view of a first example of the semiconductor radiation detector of the present embodiment, and FIG. 1B is a view taken in the direction of arrow F in FIG. A semiconductor radiation detector (crystal element assembly) 21A is a semiconductor radiation detection element in which two thin film-like electrodes A and C (conductive thin films) are formed on the surface of a semiconductor element (crystal element) S so as to be separated from each other. 211 and conductive members 22A, 22B, 23A, and 23B connected to the electrodes A and C.

  As shown in FIG. 1A, the semiconductor radiation detection element 211 is a thin film-like electrode formed on a semiconductor element S made of a plate-like semiconductor material over the entire surface of both side surfaces by sputtering or vapor deposition. Is formed. The electrode formed on one surface is an anode electrode (first conductive thin film) A, and the electrode formed on the other surface is a cathode electrode (second conductive thin film) C.

The semiconductor element S is a region that generates charges by interacting with radiation, and is formed of any single crystal such as CdTe, CdZnTe, or GaAs. The cathode electrode C and the anode electrode A are made of any material such as platinum (Pt), gold (Au), indium (In). In the semiconductor radiation detector of the first example, the semiconductor radiation detection element 211 uses, for example, CdTe for the semiconductor element S, a cathode electrode C mainly containing Pt, and an anode electrode A mainly containing In. A pn junction diode is formed.
Here, the names of the anode electrode A and the cathode electrode C are referred to as an electrode on the side where electrons generated in the semiconductor element S are collected by a high-voltage power supply applied when used as the radiation detector 21A. The electrode on the side where holes are collected is referred to as a cathode electrode C.

As shown in FIG. 1A, two plate-like conductive members 22 </ b> A and 22 </ b> B are fixed and attached to the cathode electrode C by a conductive adhesive 24 at a distance in parallel. Similarly, two plate-like conductive members 23 </ b> A and 23 </ b> B are fixed and attached to the anode electrode A in parallel by the conductive adhesive 24. The conductive members 22A, 22B, 23A, and 23B are fixed by being electrically connected to the copper foil pattern on the wiring board 113 by, for example, solder 115 in a state orthogonal to the mounting surface of the wiring board 113.
The conductive member 22A is a copper foil pattern wiring CP on the wiring board 113, the conductive member 22B is a copper foil pattern wiring 111A on the wiring board 113, and the conductive member 23A is a copper foil pattern wiring on the wiring board 113. The conductive member 23B is connected to the wiring 111B of the copper foil pattern on the wiring board 113 to the AP.

  As the conductive adhesive 24, for example, flakes (conductive particles) such as metal powder (silver) called fillers are dispersed in an organic polymer material, for example, an insulating resin binder (binder) made of an epoxy resin. Things are used. Normally, when the semiconductor radiation detection element 211 and the conductive members 22A, 22B, 23A, and 23B are bonded by the conductive adhesive 24, a high temperature of about 120 to 150 ° C. is required to cure the conductive adhesive 24. Performed by heat treatment. The conductive adhesive 24 having such a component develops conductivity when fillers come into contact with each other when the resin shrinks during the curing process of the resin.

The conductive members 22A, 22B, 23A, and 23B are flat members formed of, for example, a copper alloy. Here, phosphor bronze is used as the copper alloy.
The semiconductor radiation detector 21A corresponds to the crystal element assembly according to the first aspect.

<< Second Example of Semiconductor Radiation Detector >>
Next, a detailed configuration of the second example of the semiconductor radiation detector will be described.
The same components as those in the first example of the semiconductor radiation detector are denoted by the same reference numerals, and redundant description is omitted.
2A is a perspective view of a second example of the semiconductor radiation detector of the present embodiment, and FIG. 2B is a view taken in the direction of arrow F in FIG. The semiconductor radiation detector (crystal element assembly) 21B of this example is a thin film electrode (conductive thin film) formed on the surface of the semiconductor element (crystal element) S by sputtering, vapor deposition or the like over the entire surface of both sides. The semiconductor radiation detection element 211 is formed and resin plates 51A and 51B.

  As shown in FIG. 2A, copper foil patterns (conductive regions) 52A and 52B are electrically separated from the cathode electrode C on the surface of the plate-shaped resin plate 51A on the cathode electrode C side. Each is formed and electrically connected to the cathode electrode C by the conductive adhesive 24. Similarly, copper foil patterns 52A and 52B are formed on the surface on the anode electrode A side of the plate-shaped resin plate 51B with respect to the anode electrode A, and copper foil patterns 52A and 52B are formed. A is electrically connected to A.

As shown in FIG. 2B, the ends of the resin plates 51A and 51B are inserted into the pair of insertion holes 53a of the connector 53A. Although not understood from FIG. 2B, two fitting terminals 53b are provided in each insertion hole 53a, and are individually connected to the copper foil patterns 52A and 52B. Although the copper foil pattern 52A of the resin plate 51A is not shown, the copper foil pattern 52B of the copper foil pattern of the wiring board 113 is resinized on the wiring CP of the copper foil pattern of the wiring board 113. The copper foil pattern 52A of the plate 51B is connected to the wiring AP of the copper foil pattern of the wiring board 113, and the copper foil pattern 52B of the resin plate 51B is connected to the wiring 111B of the copper foil pattern of the wiring board 113.
The semiconductor radiation detector 21B corresponds to the crystal element assembly according to the first aspect.

<< Third example of semiconductor radiation detector >>
Next, a detailed configuration of the third example of the semiconductor radiation detector according to the present embodiment will be described.
FIG. 3A is a perspective view of a third example of the semiconductor radiation detector of the present embodiment, and FIG. 3B is a view taken in the direction of arrow F in FIG. The same components as those in the first example of the semiconductor radiation detector are denoted by the same reference numerals, and redundant description is omitted.
In the semiconductor radiation detector 21C of this example, as shown in FIG. 3, four semiconductor radiation detection elements 211 are arranged so that the cathode electrodes C and the anode electrodes A face each other, and are electrically connected to the conductive members 22A and 22B. The members 23 </ b> A and 23 </ b> B are alternately arranged between the semiconductor radiation detection elements 211 and are stacked.

A pair of conductive member 22A and conductive member 22B is interposed between the cathode electrodes C with a distance L, and a pair of conductive member 23A and conductive member 23B is interposed between the anode electrodes A with a distance L. To intervene. A pair of the conductive member 22A and the conductive member 22B is arranged at a distance L on the outermost cathode electrode C of the stacked body of the semiconductor radiation detector 21C.
The opposing surfaces of the conductive members 22A and 22B and the cathode electrode C are electrically connected and fixed by the conductive adhesive 24. Similarly, the opposing surfaces of the conductive members 23A and 23B and the anode electrode A are electrically connected and fixed by the conductive adhesive 24.

Each of the conductive members 22A and 22B and the conductive members 23A and 23B has a substantially rectangular shape, and is formed larger than the width in the height direction covering the surfaces of the electrodes A and C of the semiconductor radiation detection element 211. Further, the lateral width of the pair of conductive members 22A and 22B and the pair of conductive members 23A and 23B is formed to be larger than the lateral width covering each electrode surface of the semiconductor radiation detecting element 211 including the distance L of the gap portion. Has been.
The pair of conductive members 22A and 22B and the pair of conductive members 23A and 23B are the same as the semiconductor radiation detecting element 211 in both the width in the height direction and the width in the lateral direction including the distance L between the gaps. It does not matter even if it is large.

The thickness of the conductive members 22A, 22B, 23A, and 23B is about 10 to 100 μm, and preferably about 50 μm. Such conductive members 22A, 22B, 23A, and 23B are attached to the semiconductor radiation detection element 211, and are below the semiconductor radiation detection element 211 (wiring board 113) as shown in FIG. Projecting portions 22a and 23a depending on the side. The protruding portions 22a and 23a function as a fixing portion for attaching the semiconductor radiation detector 21C to the wiring board 113.
The protrusion 22a of the conductive member 22A is connected to the wiring CP of the copper foil pattern for the cathode electrode C provided on the wiring board 113, and the protrusion 22a of the conductive member 22B is the cathode electrode C provided on the wiring board 113. It is connected to the wiring 111A of the copper foil pattern for use. The protruding portion 23 a of the conductive member 23 A is connected to the wiring AP of the copper foil pattern for the anode electrode A provided on the wiring substrate 113, and the protruding portion 23 a of the conductive member 23 B is the anode electrode A provided on the wiring substrate 113. It is connected to the wiring 111B of the copper foil pattern for use.

Further, the semiconductor radiation detector 21C is brought into a non-contact state on the wiring board 113 by the protrusions 22a and 23a, that is, the semiconductor radiation detection element 211 has a predetermined gap between the semiconductor radiation detector 21C and the wiring board 113. It is attached.
The semiconductor radiation detector 21 </ b> C corresponds to the crystal element assembly according to claim 2.

<< Fourth Example of Semiconductor Radiation Detector >>
Next, a detailed configuration of the fourth example of the semiconductor radiation detector according to the present embodiment will be described.
FIG. 4A is a perspective view of a fourth example of the semiconductor radiation detector of the present embodiment, and FIG. 4B is a view taken in the direction of arrow F in FIG. The same components as those in the third example of the semiconductor radiation detector are denoted by the same reference numerals, and redundant description is omitted.
In the semiconductor radiation detector 21D of this example, as shown in FIG. 4, four semiconductor radiation detection elements 211 are arranged so that the cathode electrodes C and the anode electrodes A face each other, and the resin plates 51A and 51A ′ The resin plate 51B is alternately arranged between the semiconductor radiation detection elements 211 and is formed of a laminated body.
A resin plate 51A is interposed between the cathode electrodes C, and a resin plate 51B is interposed between the anode electrodes A. A resin plate 51A ′ is disposed on the outermost cathode electrode C of the stacked body of the semiconductor radiation detector 21D.
Copper foil patterns 52A and 52B are formed on the surfaces of the resin plates 51A and 51A ′ facing the cathode electrode C so as to be electrically separated and electrically connected to the cathode electrode C by the conductive adhesive 24, respectively. Yes.
Similarly, copper foil patterns 52A and 52B are formed on the surface of the resin plate 51B facing the anode electrode A so as to be electrically separated, and are electrically connected to the anode electrode A by the conductive adhesive 24, respectively. .

Each of the resin plates 51A, 51A ′, 51B has a substantially rectangular shape, and is formed larger than the width in the height direction and the width in the horizontal direction that cover each electrode surface of the semiconductor radiation detection element 211.
The resin plates 51A, 51A ′, 51B may have the same size as that of the semiconductor radiation detection element 211.

Such resin plates 51A, 51A ′, 51B are attached to the semiconductor radiation detection element 211, and are below the semiconductor radiation detection element 211 (on the wiring board 113 side) as shown in FIG. ) To project the projecting portion 51a.
As shown in FIG. 4B, the protruding portions 51a of the resin plates 51A, 51A ′, 51B are inserted into the insertion holes 53a of the connectors 53B, 53B ′ provided on the wiring board 113, respectively.

The connectors 53B and 53B ′ have essentially the same structure except that the number of the insertion holes 53a is different from three and two. Although not understood from FIG. 4B, two fitting terminals 53b are provided in the respective insertion holes 53a of the connectors 53B and 53B ′, and are individually connected to the copper foil patterns 52A and 52B. . Although not shown, the copper foil pattern 52A of the resin plates 51A and 51A ′ is on the wiring CP of the copper foil pattern of the wiring board 113, and the copper foil pattern 52B of the resin plates 51A and 51A ′ is on the wiring 111A of the copper foil pattern of the wiring board 113. The copper foil pattern 52A of the resin plate 51B is connected to the wiring AP of the copper foil pattern of the wiring board 113, and the copper foil pattern 52B of the resin plate 51B is connected to the wiring 111B of the copper foil pattern of the wiring board 113, respectively. Connected through.
Further, the semiconductor radiation detector 21D is attached in a non-contact state on the wiring board 113 by these protrusions 51a, that is, in a state where the semiconductor radiation detection element 211 has a predetermined gap between the semiconductor radiation detector 21D and the wiring board 113. .
The semiconductor radiation detector 21D corresponds to the crystal element assembly according to claim 2.

Here, the γ-ray detection characteristics of the semiconductor radiation detectors 21 of the first to fourth examples will be described. In particular, the semiconductor radiation detectors 21C and 21D of the third and fourth examples will be described.
First, the relationship between the signal rise time and the peak value curve when the thickness t (see FIGS. 3B and 4B) of the semiconductor element S is thick and thin will be described. When the reverse bias voltage of the pn junction applied between the cathode electrode C and the anode electrode A is the same value, the crest value rises (rises) faster in the semiconductor element S having a smaller thickness t, and the crest value accuracy ( Energy resolution). When the rising speed of the peak value is fast, for example, the accuracy (simultaneous counting resolution) of simultaneous measurement in the PET imaging apparatus 1 (see FIG. 7) described later is improved. The semiconductor element S having a small thickness t has a higher crest value rise rate and higher energy resolution (improves the charge collection efficiency) because the time for the electrons to reach the anode electrode A and the holes for the cathode This is because the time to reach the electrode C is shortened, that is, the charge collection time is shortened. In addition, holes that may disappear in the middle can reach the cathode electrode C without disappearing because the thickness t is thin. Incidentally, the thickness t can also be expressed as an interelectrode distance between the cathode electrode C and the anode electrode A.

The thickness (distance between electrodes) t of the semiconductor element S is preferably 0.2 mm to 2.0 mm. This is because when the thickness t exceeds 2.0 mm, the rising speed of the crest value is slowed and the maximum value of the crest value is also lowered. Even if the thickness t is increased, the moving speed of electrons and holes can be increased by increasing the reverse bias voltage and increasing the electric field strength in the thickness t direction in the semiconductor radiation detection element 211. It is possible to shorten the time for electrons and holes to reach the corresponding electrode.
However, an increase in the reverse bias voltage to be applied is not preferable because there is a problem in that the DC high-voltage power supply is increased in size and causes dielectric breakdown inside the wiring board 113. On the other hand, when the thickness is smaller than 0.2 mm, the thickness (volume) of the electrodes (cathode electrode C, anode electrode A) relatively increases. In this case, the ratio of the critical semiconductor element S that interacts with radiation decreases. That is, when the thickness t of the semiconductor element S is reduced, the thickness of the electrodes (the anode electrode A and the cathode electrode C) that do not interact with γ rays, that is, does not detect γ rays relatively increases. The proportion of the semiconductor element S that interacts with the line is relatively reduced, and as a result, the sensitivity for detecting γ-rays is lowered.

On the other hand, when the thickness t is small, the capacitance per semiconductor radiation detection element 211 increases. Since this capacitance corresponds to an input capacitance component as seen from a later-described signal processing circuit (ASIC), noise increases in the signal processing circuit as the input capacitance increases, and the energy resolution and coincidence resolution are degraded. Easy to make.
Furthermore, in order to secure a certain degree of detection sensitivity per semiconductor radiation detector 21C, 21D, the semiconductor radiation detectors 211 are arranged in parallel to effectively secure the volume of the semiconductor radiation detectors 21C, 21D. The thinner the thickness t, the greater the number of elements to be arranged in parallel. As a result, the capacitance per semiconductor radiation detector 21C, 21D increases synergistically, and performance degradation in the PET imaging apparatus 1 (see FIG. 7) (degradation in PET image contrast due to degradation in energy resolution). In addition, there is a risk of increasing inspection time and image quality degradation due to degradation of coincidence resolution. For this reason, the thickness t is preferable.

<< Electric Circuit for Crystal Element Assembly >>
Next, an electric circuit for a crystal element assembly according to the present invention will be described with reference to FIG. FIG. 5 is a schematic diagram for explaining the principle for explaining the configuration and operation of the semiconductor radiation detector and its peripheral circuits applied to the nuclear medicine diagnostic apparatus according to this embodiment.
Here, the semiconductor radiation detector 21 representatively shown in FIG. 5 may be any of the semiconductor radiation detectors 21A to 21D of the first to fourth examples described above. It is displayed in the configuration example of the radiation detector 21A.
A number of semiconductor radiation detectors 21 are incorporated in the nuclear medicine diagnosis apparatus, and one of them will be described as a representative.

A positive voltage is applied from the high voltage power supply 27 to the anode electrode A of each semiconductor radiation detection element 211 of each semiconductor radiation detector 21. At this time, the conductive member 23A is connected to the positive pole of the high-voltage power supply 27 via the wiring AP, and the conductive member 23B is connected to the positive pole of the high-voltage power supply 27 via the coil 35 and the wiring 111B.
The conductive members 23A and 23B are electrically connected to the anode electrode A by the conductive adhesive 24 described above, and the conductive members 22A and 22B are electrically connected to the cathode electrode C by the conductive adhesive 24 described above. Yes.

The cathode electrode C of each semiconductor radiation detection element 211 of the semiconductor radiation detector 21 is connected to a detection circuit 40 prepared for each semiconductor radiation detector 21. At this time, the conductive member 22A is connected to the resistor 41 and the capacitor 43 via the above-described wiring CP, and further continues to the preamplifier 45 at the rear stage of the capacitor 43. A signal processing circuit such as a fast system or a slow system is connected to the subsequent stage of the preamplifier 45, but since it is a known technique, a detailed description thereof will be omitted. The resistor 41 is grounded on the opposite side connected to the conductive member 22A.
The conductive member 22B is connected to the resistor 41 and the capacitor 43 via the wiring 111A and the coil 35 described above.
A terminal of the capacitor 43 on the semiconductor radiation detector 21 side is grounded via a resistor 41. A capacitor 46 and a resistor 47 are connected in parallel between the input side terminal and the output side terminal of the preamplifier 45.

In FIG. 5, a U-shaped yoke material 34 made of a material having a high magnetic permeability is provided near the coils 35, 35 in order to explain the operation principle of the electric circuit for the crystal element assembly of the present invention. Both ends are arranged. An excitation coil 33 is wound around the yoke material 34, and an AC power supply 31 that is on / off controlled by an on / off switch (shown as on / off SW in FIG. 5) 37 is connected to the excitation coil 33.
Here, the AC power supply 31, the exciting coil 33, the yoke material 34, the coil 35, and the on / off switch 37 correspond to the energization control means 30 described in the claims.
Incidentally, the yoke material 34 is not necessarily required for each semiconductor radiation detector 21, and an AC voltage may be induced in the coil 35, and the yoke material 34 may be disposed so that the magnetic flux generated by the excitation coil 33 passes through the coil 35. .

  When gamma rays are measured by the semiconductor radiation detector 21, when the on / off switch 37 is turned off, no AC voltage is induced in the coil 35 (first energized state), and gamma rays are incident on the semiconductor element S. When electrons and holes are generated by the photoelectric effect or the like, the electrons move to the anode electrode A and the holes move to the cathode electrode C due to the voltage applied from the high voltage power source 27, and a pulse current flows to the detection circuit 40.

  In order to maintain the electrical conductivity of the conductive adhesive 24, when the gamma ray is not measured by the radiation detector 21, when the on / off switch 37 is turned on, an alternating voltage is induced in the coil 35 (second state) On the cathode electrode C side, the coil 35 is closed to return to the coil 35 via the wiring CP, the conductive member 22A, the conductive adhesive 24, the cathode electrode C, the conductive adhesive 24, the conductive member 22B, and the wiring 111A. Alternating current flows in the circuit. Similarly, on the anode electrode A side, a closed circuit returns from the coil 35 to the coil 35 via the wiring AP, the conductive member 23A, the conductive adhesive 24, the anode electrode A, the conductive adhesive 24, the conductive member 23B, and the wiring 111B. AC current flows. It is sufficient that the current flowing through the conductive adhesive 24 by the exciting coil 33 is about several microamperes larger than several tens to several hundreds of nanoamperes flowing during γ-ray measurement.

As a result, even if an oxide film is formed on the conductive adhesive 24, it is possible to easily secure a current value sufficient to break the oxide film and improve the conductivity of the conductive adhesive 24, and to the semiconductor element S. Since the current due to this AC current does not flow, the semiconductor element S can be prevented from deteriorating.
Further, as described above, when a current larger than the current at the time of γ-ray measurement is passed through the conductive adhesive 24, it is not necessary to pass a current through the semiconductor element S. Therefore, a current that prevents an excessive current from flowing through the semiconductor element S. There is no need for a special circuit for limiting the value.

  FIG. 6 shows a perspective view of a semiconductor radiation detector of a comparative example. In the semiconductor radiation detector 231 of this comparative example, as shown in FIG. 6, four semiconductor radiation detection elements 211 are arranged so that the cathode electrodes C and the anode electrodes A face each other, and the conductive member 233 and the conductive member are arranged. 235 are alternately arranged between the semiconductor radiation detection elements 211, and are formed of a laminated body.

A conductive member 233 is interposed between the cathode electrodes C, and a conductive member 235 is interposed between the anode electrodes A. A conductive member 233 is disposed on the outermost cathode electrode C of the stacked body of the semiconductor radiation detectors 231.
Between the opposing surfaces of the conductive member 233 and the cathode electrode C are electrically connected and fixed at a plurality of locations by the conductive adhesive 24. Similarly, between the opposing surfaces of the conductive member 235 and the anode electrode A are electrically connected and fixed at a plurality of locations by the conductive adhesive 24.

  As shown in FIG. 6, the conductive members 233 and 235 are attached to the semiconductor radiation detection element 211, and as shown in FIG. 6, two conductive members 233 and 235 are suspended downward from the semiconductor radiation detection element 211 (wiring board 113 side). It has protrusions 233a and 233a and protrusions 235a and 235a. The protruding portions 233a and 235a function as a fixing portion for attaching the semiconductor radiation detector 231 to the wiring board 113. The protrusions 233a and 233a of the conductive member 233 are connected to the copper foil pattern wiring CP for the cathode electrode C provided on the wiring board 113, and the protrusions 235a and 235a of the conductive member 235 are provided on the wiring board 113. The copper foil pattern wiring AP for the anode electrode A is connected.

Even if the semiconductor radiation detector 231 having the structure as shown in the comparative example is connected by wiring instead of the semiconductor radiation detector 21 in FIG. 5, the conductive member 22A and the conductive member 22B in FIG. The conductive member 233 of FIG. 5 is replaced with a single conductive member 235 connecting the conductive member 23A and the conductive member 23B in FIG. In this case, even if the excitation coil 33 is AC-excited and AC electricity is induced in the coil 35, the current flows from the coil 35 to the coil 35 via the wiring CP, the conductive member 233, and the wiring 111 A on the cathode electrode C side. The closed circuit that returns is the main flow, and no current flows through the two conductive members 233, the conductive adhesive 24, and the anode electrode A. This is the same on the anode electrode A side.
The electrical conductivity of the conductive members 233 and 235 and the electrical conductivity of the conductive adhesive 24 are larger in the conductive members 233 and 235, especially when the conductive adhesive 24 is oxidized and its electrical conductivity is reduced. In other words, the structure of the semiconductor radiation detector 231 of the comparative example cannot be used to destroy the oxide film by energization.

《Nuclear medicine diagnostic device》
Next, a PET apparatus which is a preferred embodiment as a nuclear medicine diagnostic apparatus of the present invention will be described in detail with reference to FIGS.
As shown in FIG. 7, the PET apparatus 10 of this embodiment includes a PET imaging apparatus 1 having a measurement space 1a, a bed B that supports a subject (examinee) H, a data processing apparatus (computer or the like) 2, and the like. A display device 3 is provided. As shown in FIG. 8, the PET imaging apparatus 1 has a large number of unit substrates U arranged in the circumferential direction of the measurement space 1a. In such a PET imaging apparatus 1, the subject H is as shown in FIG. It is placed on a bed B movable in the longitudinal direction and inserted into a cylindrical measurement space 1a surrounded by the unit substrate U.

(PET imaging device)
As shown in FIG. 8, a plurality of unit substrates U of the PET imaging apparatus 1 are also arranged in the longitudinal direction of the bed B (the arrow Z direction in the drawing). As shown in FIG. 9, the unit substrate U has a radiation detection area (hereinafter referred to as a detection area) 20A and an integrated circuit area (hereinafter referred to as an ASIC (Application Specific Integrated Circuit) area) 20B. The detection region 20 </ b> A includes a plurality of semiconductor radiation detectors 21. The semiconductor radiation detector 21 detects γ rays emitted from the body of the subject H (see FIG. 1).

  The ASIC region 20B has an integrated circuit (analog ASIC 28, digital ASIC 29) for measuring the detected peak value and detection time of the γ-ray, and measures the detected peak value and detection time of the γ-ray. It has become. The integrated circuit includes a plurality of signal processing devices that process radiation detection signals.

Next, details of the PET imaging apparatus 1 will be described.
(Semiconductor radiation detector)
As the semiconductor radiation detector 21 shown in FIGS. 8 and 9, the semiconductor radiation detector 21C described in FIG. 3 or the semiconductor radiation detector 21D described in FIG. 4 is used. As shown in FIGS. 9A and 9B, such a semiconductor radiation detector 21 is formed on the wiring substrate 20 in the Y direction (radial direction of the PET imaging apparatus 1 from the detection region 20A to the ASIC region 20B, 8 rows) (see FIG. 8), 6 rows (6ch), X direction orthogonal to the Y direction (circumferential direction of the PET imaging apparatus 1, see FIG. 8), 8 rows (8ch), and the Z direction which is the thickness direction of the wiring board 20 Two surfaces (2ch) are arranged (disposed on both surfaces of the wiring board 20) in the depth direction of the PET imaging device 1 (see FIG. 8). As a result, the semiconductor radiation detector 21 is provided with a total of 48 ch on one side of the wiring board 20 and a total of 96 ch on both sides.
The method of fixing the semiconductor radiation detector 21 to the wiring board 20 is the same as the method of using the solder 115 as shown in FIG. 3 or the method of fixing the wiring board 113 using the connector as shown in FIG. .

  In the present embodiment, in order to maintain the function of detecting γ rays of each semiconductor radiation detector 21 as described above, the above-described energization control means 30 for ensuring the conductivity of the conductive adhesive 24 described in FIG. 5 is provided. ing. An exciting coil 33 is printed in the multilayer wiring substrate 20 at a position indicated by an imaginary line in FIG. 9A, and the anode electrode A side and cathode electrode of the semiconductor radiation detector 21 are adjacent thereto. A coil 35 (see FIG. 5) (not shown) is also printed on each C side.

  The cathode electrode C side of the semiconductor radiation detector 21 is connected to a detection circuit 40 (see FIG. 5). The detection circuit 40 includes a resistor 41, a capacitor 43, and an analog ASIC 28, and detects a current that flows when radiation is incident on the semiconductor radiation detector 21.

(Unit board)
As shown in FIG. 8, the unit substrate U has an annular support provided in the PET imaging apparatus 1 such that the surface on which the semiconductor radiation detector 21 is installed faces in the depth direction (Z direction) of the PET imaging apparatus 1. Installed on a member (not shown). The plurality of unit substrates U installed on the annular member are arranged in the circumferential direction and surround the measurement space 1a. And it arrange | positions so that detection area | region 20A may be located inside (measurement space 1a side) and ASIC area | region 20B may be located outside. In the present embodiment, the plurality of unit substrates U are also arranged in the depth direction of the PET imaging apparatus 1.

  The detailed structure of the unit substrate U will be described with reference to FIGS. The unit substrate U includes a detection area 20A in which a plurality of semiconductor radiation detectors 21 are installed as described above, and an ASIC area 20B. The ASIC area 20 </ b> B includes a resistor 41, a capacitor 43, an analog ASIC 28, and a digital ASIC 29. The γ-ray detection signal output from each semiconductor radiation detector 21 is supplied from the detection region 20A side to the ASIC region 20B side by a multilayer wiring (not shown) in the wiring board 20.

In the ASIC region 20B, as shown in FIGS. 8A and 8B, one digital ASIC 29 is installed on one side and four analog ASICs 28 are arranged on both sides. On both surfaces of the ASIC region 20B, resistors 41 and capacitors 43 are installed in a number corresponding to the number of semiconductor radiation detectors 21.
A plurality of connection wirings (not shown) for electrically connecting the resistor 41, the capacitor 43, the analog ASIC 28, and the digital ASIC 29 are provided in the ASIC region 20B. The analog ASIC 28 means an application specific integrated circuit (ASIC) that processes an analog signal (γ-ray detection signal) output from the semiconductor radiation detector 21, and is an LSI (Large Scale Integrated Circuit) of the LSI. It is a kind. The analog ASIC 28 is provided with a signal processing circuit for each semiconductor radiation detector 21. These signal processing circuits are adapted to obtain a γ-ray peak value by inputting a γ-ray detection signal (radiation detection signal) output from one corresponding semiconductor radiation detector 21.
Incidentally, the analog ASIC 28 is a fast and slow signal processing circuit including the preamplifier 45 in FIG.

  The shorter the circuit length and the length (distance) of the wiring for transmitting the γ-ray detection signal, the smaller the capacitance, which is preferable because the influence of noise is small. Therefore, in this embodiment, the semiconductor radiation detector 21, the capacitor 43, the resistor 41, the analog ASIC 28, and the digital ASIC 29 are arranged in this order on the unit substrate U from the central axis toward the outside in the radial direction of the PET imaging apparatus 1. It is arranged. With this configuration, the length (distance) of the wiring that transmits the weak γ-ray detection signal output from the semiconductor radiation detector 21 to the amplifier of the analog ASIC 28 can be shortened. For this reason, the influence of noise on the detection signal of γ rays is reduced.

  As described above, the detection signal output from the semiconductor radiation detector 21 is output from the cathode electrode C side to the conductive members 22A and 22B via the conductive adhesive 24. Since the detection signal is a high-frequency signal of 10 MHz or more, if there is a portion with poor conductivity in the conductive adhesive 24, distortion occurs when the detection signal passes through this portion, which is a cause of noise. It appears as degradation of characteristics equivalent to the increase. For this reason, it is necessary to maintain good electrical conductivity of the conductive adhesive 24, and as a means for that purpose, an operation described below is required at the time of startup.

First, before starting the PET imaging apparatus 1, by operating an operation panel (not shown) or the like, the on / off switch 37 (see FIG. 5) is turned on to supply the AC power supply 31 (see FIG. 5). An excitation coil 33 (see FIG. 9) built in 20 is excited by an AC power supply 31, and an AC voltage is induced in the coil 35 (see FIG. 5). As a result, on the cathode electrode C side, the coil 35 shown in FIG. 5 is passed through the wiring CP, the conductive member 22A, the conductive adhesive 24, the cathode electrode C, the conductive adhesive 24, the conductive member 22B, and the wiring 111A. A maximum of several microamperes of alternating current flows in the closed circuit. Similarly, on the anode electrode A side, a closed circuit returns from the coil 35 to the coil 35 via the wiring AP, the conductive member 23A, the conductive adhesive 24, the anode electrode A, the conductive adhesive 24, the conductive member 23B, and the wiring 111B. An alternating current of up to several microamperes flows.
Note that these operations may be automatically performed according to a program (not shown) set in advance, for example, before normal startup or every other week.

  When a current larger than several tens to several hundreds of nanoamperes flowing at the time of γ-ray measurement flows to the bonding structure formed by the conductive adhesive 24 of the semiconductor radiation detector 21, the conductive particles contained in the conductive adhesive 24 are The oxide film formed on the conductive adhesive 24 is re-bonded or broken, and the conductivity is stabilized.

The conductivity of the conductive adhesive 24 significantly affects the characteristics of the semiconductor radiation detector 21, that is, the energy resolution and time accuracy, and causes deterioration. Greatly contributes to securing characteristics. In the present embodiment, the conduction control means 30 (see FIG. 5) can maintain good electrical conductivity. Therefore, a structure in which the semiconductor radiation detection element 211 and the conductive members 22A, 22B, 23A, and 23B are stacked, or a semiconductor Even when the radiation detection element 211 and the resin plates 51A, 51A ′, 51B are stacked, the deterioration of characteristics is reduced.
In addition, since the electrical conductivity can be ensured, the semiconductor radiation detection element 211 can be formed to be thinner and stacked, and the performance and sensitivity of the semiconductor radiation detector 21 can be improved at the same time.

Below, the effect acquired in this embodiment is demonstrated.
(1) In the semiconductor radiation detectors 21A to 21D of the first to fourth examples, at least two conductive members 22A and 22B (or conductive members 23A) that are electrically separated from the cathode electrode C and the anode electrode A are used. , 23B) or copper foil patterns 52A, 52B are connected by a conductive adhesive 24. Therefore, on the cathode electrode C side, from the coil 35, the wiring CP, the conductive member 22A (or the copper foil pattern 52A), the conductive adhesive 24, the cathode electrode C, the conductive adhesive 24, and the conductive member 22B (or the copper foil pattern 52B). ), The closed circuit returning to the coil 35 via the wiring 111A is similarly connected on the anode electrode A side from the coil 35 to the wiring AP, the conductive member 23A (or the copper foil pattern 52A), the conductive adhesive 24, the anode electrode A, It is possible to configure a closed circuit that returns to the coil 35 via the conductive adhesive 24, the conductive member 23B (or the copper foil pattern 52B), and the wiring 111B.
(2) Since the power supply control means 30 is provided to allow at least a current larger than the current due to the charge generated when γ-rays (radiation) is incident to flow through the conductive adhesive 24, the power supply control means The conductive particles of the conductive adhesive 24 can be recombined or the oxide film formed on the conductive adhesive 24 can be broken by the large current flowing through the conductive adhesive 24. Can be stabilized.
(3) A nuclear medicine diagnostic apparatus excellent in energy resolution can be obtained with the minimum additional configuration that the wiring for supplying the AC power supply 31 and the exciting coil 33 are provided on the wiring board 20 of the unit board U.
(4) By energizing the energization control means 30 before applying a voltage for measurement, it is possible to effectively suppress the failure of the semiconductor radiation detector 21 and maintain the radiation detection performance over a long period of time. be able to.
(5) Since the conductivity of the conductive adhesive 24 can be maintained, the semiconductor radiation detection element 211 can be formed to be thinner and stacked, and both the performance and sensitivity of the semiconductor radiation detector 21 can be improved. It becomes possible to make it. In the PET imaging apparatus 1, it is necessary to efficiently capture 511 keV γ-rays. For this purpose, the semiconductor radiation detection element 211 must be thickened. However, if the semiconductor radiation detection element 211 is thickened, the moving distance of electrons and holes becomes long, so that the energy resolution and the recognition accuracy of the incident time are deteriorated. If a large number of thin semiconductor radiation detection elements 211 can be stacked, the movement distance of electrons and holes can be shortened, so that the energy resolution and the recognition accuracy of the incident time can be improved. There is an advantage that the volume of the container 21 can be increased, and the performance of the PET imaging apparatus 1 can be improved.
(6) Since the semiconductor radiation detector 21 is disposed so that the cathode electrodes C or the anode electrodes A of the semiconductor radiation detection elements 211 adjacent to each other face each other, the conductive members 22A, 22B, 23A, and 23B are shared. be able to. In addition, since the conductivity of the conductive adhesive 24 can be maintained, a reverse bias voltage for collecting charges can be applied between the cathode electrode C and the anode electrode A in a stable state.
Further, it is not necessary to arrange an electrical insulating material between the semiconductor radiation detection elements 211, and a dense arrangement of the semiconductor radiation detection elements 211 can be realized. Thereby, the sensitivity is improved and the inspection time can be shortened.

(Modification of energization control means)
The energization control means 30 in the PET apparatus 10 of the present embodiment is provided with the exciting coil 33 and the coil 35 on the unit substrate U, but is not limited to this.
For example, the wiring board 20 of the unit board U is provided with only the coil 35, and a detachable exciter that excites an excitation coil 33 wound around a yoke material 34 with an AC power supply 31 is provided separately. An excitation device may be attached to the measurement space 1a before the γ-ray measurement, and the coil 35 may be excited by the magnetic flux from the yoke material 34.

  Further, if the magnetic flux density in the direction perpendicular to the paper surface in FIG. 5 is changed, the wiring CP, the conductive member 22A, the conductive adhesive 24, the cathode electrode C, and the conductive adhesive are provided on the cathode electrode C side even without the coil 35. 24, an alternating current flows in a closed circuit that returns to the wiring CP through the conductive member 22B and the wiring 111A. Similarly, on the anode electrode A side, an alternating current is generated in a closed circuit that returns to the wiring AP through the wiring AP, the conductive member 23A, the conductive adhesive 24, the anode electrode A, the conductive adhesive 24, the conductive member 23B, and the wiring 111B. Flowing. Therefore, in such a case, the energization control means 30 need only be an excitation device that supplies a magnetic flux density that changes from the outside to the closed circuit, and the wiring board 20 is further simplified.

  In this embodiment, the PET imaging apparatus 1 has been described as an example of a nuclear medicine diagnosis apparatus. However, an SPECT apparatus, a gamma camera, an X-ray CT apparatus, and an X-ray CT-PET apparatus that combines an X-ray CT apparatus and a PET apparatus are used. Is also applicable.

<< Other Embodiments >>
In the present embodiment, the semiconductor radiation detector 21 using a semiconductor crystal suitable for the PET apparatus 10 has been described as an example of a crystal element assembly using a crystal element that is a functional element, but is not limited thereto. .

For example, when applied to a solar cell, single crystal silicon, polycrystalline silicon or the like is used as a crystal element as a functional element, a conductive thin film is formed as an electrode on them, and two or more electrically separated electrodes connected to an external circuit The conductive member thus formed may be connected to a conductive thin film as an electrode using a conductive adhesive. As the conductive adhesive in this case, a glass binder may be used instead of a resin binder, and flakes (conductive particles) of a conductive substance such as silver called filler may be blended. Thus, by providing a plurality of conductive members on one electrode of a solar cell element, a circuit of conductive member-conductive adhesive-electrode-conductive adhesive-conductive member can be formed. Thus, it is possible to configure an energization control means that maintains the continuity of the current.
In addition, the crystal element which is a functional element in the present invention includes amorphous (amorphous).
Furthermore, the crystal element assembly of the present invention can also be applied to piezoelectric elements such as Rochelle salt.

(A) is a perspective view of the 1st example of the semiconductor radiation detector of embodiment of this invention, (b) is F arrow line view in (a). (A) is a perspective view of the 2nd example of the semiconductor radiation detector of this embodiment, (b) is F arrow figure in (a). (A) is a perspective view of the 3rd example of the semiconductor radiation detector of this embodiment, (b) is F arrow figure in (a). (A) is a perspective view of the 4th example of the semiconductor radiation detector of this embodiment, (b) is F arrow figure in (a). It is the circuit diagram which showed the characteristic part of this embodiment. It is a perspective view of the semiconductor radiation detector of a comparative example. It is the perspective view which showed typically the structure of the PET apparatus suitable as a nuclear medicine diagnostic apparatus of this embodiment. It is the perspective view which showed the PET imaging device typically. (A) is a front view of the unit board | substrate used for the PET imaging device shown in FIG. 8, (b) is a side view of a unit board | substrate similarly.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 PET imaging device 1a Measurement space 2 Data processing device 3 Display device 10 PET device 20 Wiring board 20A Detector region 20B ASIC region 21, 21A, 21B, 21C, 21D Semiconductor radiation detector (crystal element assembly)
22A, 22B, 23A, 23B Conductive member 24 Conductive adhesive 27 High-voltage power supply 28 Analog ASIC
29 Digital ASIC
30 Energization control means 31 AC power supply (energization control means)
33 Excitation coil (energization control means)
34 Yoke material (energization control means)
35 Coil (energization control means)
37 ON / OFF switch (energization control means)
40 Detection Circuit 51A, 51B Resin Plate 52A, 52B Copper Foil Pattern (Conductive Area)
53A, 53B Connector 53a Insertion hole 53b Fitting terminal 111A, 111B Wiring 113 Wiring board 115
211 Semiconductor radiation detector (unit)
A Anode electrode (first conductive thin film)
AP, CP wiring C cathode electrode (second conductive thin film)
H Subject S Semiconductor element (crystal element)
U unit board

Claims (10)

  At least two conductive members for electrical connection to the outside are provided for each of the first and second conductive thin films formed on the surface of the crystal element which is a functional element, and the conductive particles and An adhesive structure in which the at least two conductive members and the first or second conductive thin film are bonded to the first and second conductive thin films by a conductive adhesive composed of a binder. A crystal element assembly comprising: A unit in which first and second conductive thin films are formed on two opposing surfaces of a crystal element which is a functional element;
Conductivity for electrically connecting the unit to the first conductive thin film and the second conductive thin film so that the first and second conductive thin films are electrically connected to the outside. Provide at least two members, laminate,
The at least two conductive members and the first or second conductive thin film are bonded to the conductive member and the first and second conductive thin films by a conductive adhesive composed of conductive particles and a binder. A crystal element assembly having an adhesion structure formed by bonding. Using the crystal element assembly according to claim 1 or 2,
The first conductive state of the conductive member in which the first conductive thin film, the crystal element, and the second conductive thin film serve as a current path, and one conduction of the first and second conductive thin films. An electric circuit for a crystal element assembly, characterized by combining a current-carrying control means capable of switching between a second current-carrying state of the conductive member between the adhesive structures on the conductive thin film. The energization control means includes
The electric current generated when the crystal element acts as a functional element in the first energized state is taken out as an electric signal from the conductive member connected to the first and second conductive thin films,
In the second energized state, a current larger than an electric signal current caused by electric charges generated when acting as the functional element is caused to flow between the adhesive structures on one conductive thin film. The electrical circuit for a crystal element assembly according to claim 3. The crystal element is made of a semiconductor crystal,
5. The electric circuit for a crystal element assembly according to claim 4, wherein the crystal element assembly constitutes a semiconductor radiation detector.   The electric circuit for a crystal element assembly according to any one of claims 3 to 5, wherein the conductive member is a plate made of a metal material.   The conductive member is at least two conductive regions formed on the surface of a plate made of an electrically insulating material, and the at least two conductive regions are formed to be electrically separated from each other. The electrical circuit for a crystal element assembly according to any one of claims 3 to 5, wherein:   The electric circuit for a crystal element assembly according to claim 7, wherein the plate is made of a resin material.   A nuclear medicine diagnosis apparatus using the electrical circuit for a crystal element assembly according to any one of claims 3 to 8. Using the crystal element assembly according to claim 1 or 2,
The first conductive state of the conductive member in which the first conductive thin film, the crystal element, and the second conductive thin film serve as a current path, and one conduction of the first and second conductive thin films. A conduction control method in an electric circuit for a crystal element assembly, in which a conduction control means capable of switching between a second conduction state of the conductive member in which a current path is formed between the bonding structures on the conductive thin film is combined. ,
The energization control means includes
The electric current generated when the crystal element acts as a functional element in the first energized state is taken out as an electric signal from the conductive member connected to the first and second conductive thin films,
The second energized state is controlled so that a current larger than an electric signal current caused by electric charges generated when acting as the functional element flows between the adhesive structures on one conductive thin film. A method for controlling energization in an electric circuit for a crystal element assembly.

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