JP5546806B2 - Nuclear medicine imaging equipment - Google Patents

Description

The present invention relates to a nuclear medicine imaging apparatus.

  Conventionally, nuclear medicine imaging apparatuses such as a gamma camera, a single photon emission CT apparatus (SPECT apparatus), and a positron emission CT apparatus (PET apparatus) are medical image diagnosis apparatuses capable of performing functional diagnosis on a living tissue of a subject. It is widely used in today's medical field (see, for example, Patent Document 1).

  Specifically, a nuclear medicine imaging device detects gamma rays emitted from an isotope or a labeled compound selectively taken into a living tissue by a detector, and a nuclear medicine image obtained by imaging a dose distribution of the detected gamma rays. Is a device for reconfiguring.

  In recent years, a device (for example, a PET-CT device or a SPECT-CT) in which a nuclear medicine imaging device and an X-ray CT (Computed Tomography) device for imaging morphological information in a living tissue of a subject are integrated. Devices) have been put into practical use.

  Here, a detector of a nuclear medicine imaging apparatus generally includes a scintillator that converts incident radiation such as gamma rays into light having a peak in the ultraviolet region, and light emission (photoelectrons) from the scintillator is multiplied to generate electricity. A photomultiplier tube (PMT) that converts the signal into a signal is used (see, for example, Non-Patent Document 1).

  Specifically, the PMT includes a photocathode that receives scintillation light and generates photoelectrons, a multistage dynode that provides an electric field that accelerates the generated photoelectrons, and an anode that is an outlet for electrons. Electrons emitted from the photocathode due to the photoelectric effect are accelerated toward the dynode, collide with the surface of the dynode, and knock out a plurality of electrons. By repeating this phenomenon over multiple dynodes, the number of electrons is avalancheally increased, and the number of electrons at the anode reaches about 1 million. In such an example, the gain factor of the PMT is 1 million times. In addition, a voltage of 1000 volts or more is usually applied between the dynode and the anode for amplification using the avalanche phenomenon.

  However, in the PMT, since the movement of electrons is affected by the magnetic field, it is necessary to provide a magnetic shield around the PMT. In addition, since the PMT is affected by a magnetic field, an apparatus that integrates an MRI apparatus for imaging morphological information in a biological tissue of a subject using a nuclear magnetic resonance phenomenon and a nuclear medicine imaging apparatus uses scintillation light. The effect of the magnetic field could not be realized unless a special structure was adopted such as guiding it to the outside with an optical fiber and inputting it to the PMT.

Here, a silicon photomultiplier (SiPM) using an avalanche photodiode (APD) as a semiconductor element array is known as a photomultiplier having a gain factor equivalent to that of the PMT, and will be used in the near future. It is considered to replace PMT in a wide range of fields. A reverse voltage is applied to the SiPM APD element. When one scintillation photon is incident on the pixel, the pixel generates an electric pulse of 10 ⁵ to 10 ⁶ electrons. Here, unlike the aforementioned PMT, the SiPM does not require an amplifier circuit.

  Factors for which SiPM is expected to be widespread include the fact that it has a gain factor equivalent to that of PMT as described above, and that “the quantum effect is high compared to PMT”, and “because it is not affected by a magnetic field, In addition, an apparatus in which an MRI apparatus and a nuclear medicine imaging apparatus are integrated can be put into practical use "," pixel detection is possible "," required applied voltage is, for example, , 17V to 70V, which is lower than PMT and easy to handle, ”“ because it can be manufactured by a CMOS process, low-cost mass production is expected ”, and the like. Due to these factors, SiPM is expected to be used not only in the detector of nuclear medicine imaging apparatus but also in the detector of X-ray CT apparatus that detects X-rays.

  By the way, the gain factor of SiPM has sensitive temperature dependence and applied voltage dependence. In particular, when SiPM is used as a detector of a nuclear medicine imaging apparatus or an X-ray CT apparatus, a control mechanism for stabilizing the gain factor even if the temperature fluctuates by controlling the gain factor change due to the temperature change is indispensable. Become.

  As such a control mechanism, there is a temperature control mechanism for controlling the temperature of the entire apparatus in which the SiPM is incorporated to be constant. That is, by incorporating a temperature control mechanism in the apparatus, the temperature around SiPM and SiPM becomes constant, and the gain factor of SiPM becomes stable.

JP 2007-107995 A

Edited by Japan Imaging and Medical Systems Industry Association “Medical Image / Radiological Equipment Handbook” Meiko Art Printing Co., Ltd. 2001, p. 183-184

  By the way, the above-described conventional technology requires a complicated device structure in order to realize the temperature control mechanism, and the gain factor of the detector using SiPM cannot be easily stabilized corresponding to the temperature change. There was a problem.

  In the above description, the SiPM that multiplies photoelectrons by applying a reverse voltage to the APD has been described as having a problem that the gain factor cannot be easily stabilized in response to a temperature change. However, even in a device that multiplies photoelectrons by applying a forward voltage to the APD, there is a problem that the gain factor cannot be easily stabilized corresponding to a temperature change in the same way as SiPM due to the characteristics of the APD.

Accordingly, the present invention has been made to solve the above-described problems of the prior art, and provides a nuclear medicine imaging apparatus capable of easily stabilizing a gain factor in response to a temperature change. Objective.

In order to solve the above-described problems and achieve the object, the present invention provides a scintillator that converts gamma rays into light and a plurality of amplifying elements that multiply the light converted by the scintillator as electric signals in a two-dimensional array. The SiPM having the amplifying element group arranged and the applied voltage applied to the amplifying element are controlled to be divided so as to increase according to the increase in the ambient temperature, and decrease according to the decrease in the ambient temperature. In order to control the voltage division in such a manner, a semiconductor resistance that decreases as the ambient temperature increases and increases as the ambient temperature decreases is connected in series with the semiconductor resistance. , the resistance value is composed as a fixed stationary resistance to changes in environmental temperature, and a voltage control mechanism is connected in parallel with the amplifying element to which the fixed resistance is the partial pressure control of the target, the voltage control mechanism For each amplifier element comprising a partial pressure control of the target, or, characterized in that it is installed in each amplifying element group of a plurality of amplifying elements as a partial pressure control of the subject.

According to the present invention, it is possible to easily stabilize the gain factor corresponding to a temperature change.

FIG. 1 is a diagram for explaining a configuration of a SPECT apparatus according to the present embodiment. FIG. 2 is a diagram for explaining the configuration of the gamma camera in the present embodiment. FIG. 3 is a diagram for explaining the characteristics of the gain factor of SiPM. FIG. 4 is a diagram for explaining the configuration of the SiPM in this embodiment. FIG. 5 is a diagram for explaining the characteristics of the gain factor of SiPM in this embodiment. FIG. 6 is a diagram for explaining a modification of the present embodiment.

  Exemplary embodiments of a photomultiplier according to the present invention will be described below in detail with reference to the accompanying drawings. Hereinafter, a case where the photomultiplier according to the present invention is mounted on a single photon emission CT apparatus (SPECT apparatus) which is a nuclear medicine imaging apparatus will be described as an example.

  First, the configuration of the SPECT apparatus in this embodiment will be described. FIG. 1 is a diagram for explaining a configuration of a SPECT apparatus according to the present embodiment. As shown in FIG. 1, the SPECT apparatus according to the present embodiment includes a gantry device 10 and a console device 20.

  The gantry device 10 is a device that collects projection data by detecting gamma rays emitted from a radiopharmaceutical administered to the subject P and selectively taken into the living tissue of the subject P. The bed 12 includes a bed driving unit 13, a gamma camera 14, and a camera driving unit 15. As shown in FIG. 1, the gantry device 10 has a cavity serving as a photographing port.

  The top plate 11 is a bed on which the subject P lies, and is placed on the bed 12. The couch driving unit 13 moves the subject P into the imaging port of the gantry device 10 by moving the couch 12 under the control of the couch controller 23 described later.

  The gamma camera 14 two-dimensionally detects the intensity distribution of gamma rays emitted from a radiopharmaceutical nuclide (RI: Radio Isotope) that is selectively taken into the living tissue of the subject P, and detects the detected two-dimensional gamma ray intensity. For example, the distribution data is an apparatus that generates projection data by performing amplification processing and A / D conversion processing, and transmits the generated projection data to a data collection unit 25 described later.

  The camera drive unit 15 is a device that moves the gamma camera 14 under the control of a camera control unit 24 described later. For example, the camera drive unit 15 drives the gamma camera 14 along the imaging port of the gantry device 10, and thereby the gamma camera 14 rotates around the subject P to obtain projection data in a 360 degree direction. Generate.

  The console device 20 accepts the operation of the SPECT device by the operator, and at the same time, a nuclear medicine image (SPECT) that is a tomographic image depicting the distribution of the radiopharmaceutical administered to the subject P from the projection data collected by the gantry device 10. Device).

  Specifically, as illustrated in FIG. 1, the console device 20 includes an input unit 21, a display unit 22, a bed control unit 23, a camera control unit 24, a data collection unit 25, and an image reconstruction unit 26. The data storage unit 27 and the system control unit 28 are connected to each unit of the console device 20 via an internal bus.

  The input unit 21 includes a mouse, a keyboard, and the like that are used by the operator of the SPECT apparatus to input various instructions and various settings, and transfers instructions and setting information received from the operator to the system control unit 28.

  The display unit 22 is a monitor that is referred to by the operator, and displays a SPECT image or the like to the operator under the control of the system control unit 28 or various instructions and various settings from the operator via the input unit 21. For example, a GUI (Graphical User Interface) is received.

  The data collection unit 25 collects the projection data transmitted from the gamma camera 14 and performs correction processing such as logarithmic conversion processing, offset correction, and sensitivity correction on each of the collected projection data to obtain corrected projection data. The generated corrected projection data is stored in the data storage unit 27.

  The image reconstruction unit 26 reads the corrected projection data from the data storage unit 27, and performs the back projection process on the read corrected projection data (for example, 360 degree direction corrected projection data), thereby performing a SPECT image. Reconfigure. Then, the image reconstruction unit 26 stores the reconstructed SPECT image in the data storage unit 27.

  The system control unit 28 performs overall control of the SPECT device by controlling the operations of the gantry device 10 and the console device 20. Specifically, the system control unit 28 controls the bed control unit 23 and the camera control unit 24 to execute the projection data collection process in the gantry device 10. Further, the system control unit 28 controls the entire image processing in the console device 20 by controlling the correction processing of the data collection unit 25 and the reconstruction processing of the image reconstruction unit 26. Further, the system control unit 28 controls the display unit 22 to display the SPECT image stored in the data storage unit 27.

  Here, the gamma camera 14 in the present embodiment is not a photomultiplier tube (PMT) that has been generally used in the past, but multiplies input light as an electrical signal with a gain factor equivalent to that of the PMT. It is comprised by the photomultiplier which performs. Specifically, the gamma camera 14 in the present embodiment is configured by a silicon photomultiplier (SiPM) in which an avalanche photodiode (APD) that is not affected by a magnetic field is used as a semiconductor element array. The Hereinafter, the configuration of the gamma camera in the present embodiment will be described with reference to FIG. FIG. 2 is a diagram for explaining the configuration of the gamma camera in the present embodiment.

  Specifically, as shown in FIG. 2A, the gamma camera 14 in this embodiment includes a scintillator 14a and a SiPM 14b, and the scintillator 14a receives gamma rays emitted from the internal tissue of the subject P. The light is converted into light having a peak in the ultraviolet region, and the SiPM 14b multiplies the light input from the scintillator 14a as an electrical signal by the APD at a gain rate equivalent to that of the PMT.

  Thereby, the gamma camera 14 in the present embodiment generates projection data from the detected two-dimensional gamma ray intensity distribution data, and transmits the generated projection data to the data collection unit 25 as shown in FIG. To do.

  Here, the SiPM 14b is configured, for example, by arranging a plurality of APDs 14c in a two-dimensional array as shown in FIG.

  However, the gain factor of the SiPM 14b has sensitive temperature dependency and applied voltage dependency. Specifically, as shown in FIG. 3A, the gain factor of the SiPM 14b has a characteristic of decreasing as the temperature increases. Furthermore, as shown in FIG. 3A, the gain factor of the SiPM 14b has a characteristic of increasing as the applied voltage increases. FIG. 3 is a diagram for explaining the characteristics of the gain factor of SiPM.

  Therefore, the gamma camera 14 according to the present embodiment has a main feature that the gain factor can be easily stabilized corresponding to the temperature change by installing the voltage control mechanism described below in the SiPM 14b. . This main feature will be described with reference to FIGS. FIG. 4 is a diagram for explaining the configuration of the SiPM in this example, and FIG. 5 is a diagram for explaining the characteristics of the gain factor of the SiPM in this example.

  Here, as described in FIG. 2B, the SiPM 14b in the present embodiment has a plurality of APDs 14c arranged in a two-dimensional array, but in this embodiment, a plurality of APDs 14c are provided for each predetermined number. A plurality of divided APD arrays 14d are provided, and a voltage control mechanism is installed in each APD array 14d. For the APD array 14d, as shown in FIG. 4, the applied voltage applied to each APD 14c constituting the APD array 14d is a voltage in a direction opposite to the direction of the voltage generated by each APD 14d ( Reverse voltage).

  Then, the SiPM 14b performs voltage division control so that the voltage applied to the APD array 14d increases as the ambient temperature increases, and also performs voltage division control so as to decrease as the ambient temperature decreases. Is provided.

  Specifically, as shown in FIG. 4, the SiPM 14 b has a semiconductor resistance 14 e that decreases in resistance value as the ambient temperature increases and increases in resistance value as the ambient temperature decreases, and a semiconductor resistance 14 e. Are connected in series with each APD 14c of the APD array 14d to be subjected to voltage division control. A voltage control mechanism is installed for each APD array 14d to be subjected to voltage division control.

  Here, as shown in FIG. 4, the resistance value of the fixed resistor 14f is “R”, the resistance value of the semiconductor resistor 14e that varies depending on the temperature (t) is “r (t)”, and the fixed resistor 14f The voltage applied to the semiconductor resistor 14e is “V”, and the voltage applied to the APD array 14d is “Vinv”.

  In such a case, “Vinv” is expressed as “V × (R / (R + r (t)))”. As described above, the resistance value of “r (t)” decreases as the ambient temperature increases, and the resistance value increases as the ambient temperature decreases. Therefore, “Vinv” The temperature rises as the temperature rises, and falls as the ambient temperature in the SiPM 14b decreases.

  As a result, the gain factor of the SiPM 14b in the present embodiment has a characteristic that becomes a stable value corresponding to a temperature change, as shown in FIG.

  As described above, in this embodiment, when the gamma camera 14 installed in the SPECT apparatus is configured using the SiPM 14b having a plurality of APDs 14c arranged in a two-dimensional array, all the APDs 14c are arranged from a predetermined number of APDs 14c. The APD array 14d is divided and a voltage control mechanism is installed for each APD array 14d.

  In other words, the voltage control mechanism is connected in series with the semiconductor resistor 14e, the resistance value of which decreases as the ambient temperature increases and the resistance value increases as the ambient temperature decreases, and the semiconductor resistor 14e. Is formed of a fixed resistor 14f that is fixed with respect to changes in the ambient temperature, and the fixed resistor 14f is connected in parallel to each APD 14c of the APD array 14d that is subject to voltage division control.

  Therefore, when the ambient temperature rises, the resistance value of the semiconductor resistor 14e decreases, so that the voltage applied to the APD array 14d increases, and as a result, the gain factor that decreases as the ambient temperature increases increases due to the increase in applied voltage. As described above, the gain factor can be easily stabilized corresponding to the temperature change.

  In addition, by installing a voltage control mechanism independently for each APD array 14d, even if a temperature change occurs locally inside the gamma camera 14, only the location where the temperature change occurs is locally Since fluctuations due to temperature dependence of the gain factor can be mitigated, the need for a complicated mechanism for temperature management of the entire SPECT apparatus can be eliminated, and the cost required for the SPECT apparatus can be greatly reduced. It becomes.

  In addition, the degree of freedom in designing the gamma camera 14 can be improved due to the independence in installing the voltage control mechanism. Further, since the SiPM 14b can be manufactured by a CMOS process, a large number of APDs 14c can be mounted on the gamma camera 14, and the effective field of view of the SPECT image can be improved. Further, since the SiPM 14b can be manufactured by a CMOS process, the failed SiPM 14b can be easily replaced. Furthermore, since each SiPM 14b has an independent voltage control mechanism, the influence of replacement is less likely to affect the entire system, adjustment costs after replacement can be greatly reduced, and the cost required to maintain the gamma camera 14 is also reduced. It becomes possible to reduce. In addition, by using SiPM 14b that is not affected by magnetism, it is possible to realize an apparatus in which a SPECT apparatus or a PET apparatus is integrated with an MRI apparatus.

  The present invention may be implemented in various different forms other than the above-described embodiments. Therefore, in the following, a modification of the present embodiment will be described with reference to FIG. In addition, FIG. 6 is a figure for demonstrating the modification in a present Example.

  As described in the above embodiment, the voltage control mechanism including the semiconductor resistor 14e and the fixed resistor 14f is installed with respect to the APD array 14d as shown in FIG. The number of APDs 14c constituting the APD array 14d is arbitrarily set by the designer of the gamma camera 14.

  For example, the APD array 14d is composed of 9 APDs 14c (see (1) in FIG. 6A) or 25 APDs 14c (see (2) in FIG. 6A). ). Alternatively, the voltage control mechanism including the semiconductor resistor 14e and the fixed resistor 14f may be provided for each APD 14c (see (3) in FIG. 6A).

  Further, in the above embodiment, by applying a reverse voltage to the APD 14c, the device (SiPM 14b) in which the voltage control mechanism is installed in the photomultiplier used in the Geiger mode is mounted on the SPECT device as the gamma camera 14. Explained when to do. However, the present invention is not limited to this, and as shown in FIG. 6B, a voltage control mechanism is installed in a photomultiplier that doubles light by applying a forward voltage to the APD 14c. For example, the apparatus may be mounted on a nondestructive inspection apparatus using radiation such as X-rays.

  In the above embodiment, the case where the SiPM 14 provided with the voltage control mechanism is mounted on the SPECT apparatus has been described. However, the present invention is not limited to this, and the SiPM 14b provided with the voltage control mechanism is installed. It may be mounted as a gamma ray detector of a nuclear medicine imaging apparatus other than a SPECT apparatus such as a gamma camera or a positron emission CT apparatus (PET apparatus) or an X-ray detector of an X-ray CT apparatus.

As described above, the nuclear medicine imaging apparatus according to the present invention is useful when the light is multiplied as an electric signal by the amplifying element, and is particularly suitable for easily stabilizing the gain factor corresponding to a temperature change. .

DESCRIPTION OF SYMBOLS 10 Stand apparatus 11 Top plate 12 Sleeper 13 Sleeper drive part 14 Gamma camera 14a Scintillator 14b SiPM (silicon photomultiplier)
14c APD (avalanche photodiode)
14d APD array 14e Semiconductor resistance 14f Fixed resistance 15 Camera drive unit 20 Console device 21 Input unit 22 Display unit 23 Bed control unit 24 Camera control unit 25 Data collection unit 26 Image reconstruction unit 27 Data storage unit 28 System control unit 28 System control unit

Claims (3)

A scintillator that converts gamma rays into light;
SiPM having an amplification element group in which a plurality of amplification elements that multiply the light converted by the scintillator as an electric signal are arranged in a two-dimensional array;
In order to perform voltage division control so that an applied voltage applied to the amplification element increases in accordance with an increase in ambient temperature, and to perform voltage division control so as to decrease in accordance with a decrease in the ambient temperature, A semiconductor resistor that decreases in resistance with an increase in temperature and increases in resistance with a decrease in the ambient temperature, and is connected in series with the semiconductor resistor. A fixed resistance that is fixed, and the fixed resistance includes a voltage control mechanism that is connected in parallel with an amplification element that is a target of voltage division control,
The nuclear medicine imaging apparatus, wherein the voltage control mechanism is installed for each amplification element that is a target of voltage division control or for each amplification element group that is a plurality of amplification elements that are targets of voltage division control .   The nuclear medicine imaging apparatus according to claim 1, wherein the applied voltage is a voltage in a direction opposite to a direction of a voltage generated by the amplifying element.   The nuclear medicine imaging apparatus according to claim 1, wherein the applied voltage is a forward voltage with respect to a direction of a voltage generated by the amplifying element.

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