KR101096977B1 - Positron emission tomography devices and scintillator detectors for the same - Google Patents

Abstract

The present invention includes sensing cells for sensing a flash signal from the crystal layers, and crystal layers sequentially formed on one surface of the photomultiplier tube and the photomultiplier tube where the sensing cells are located to amplify and output photoelectrons emitted from the sensing cells. A scintillation detector comprising at least three or more multilayer crystal layers, wherein the crystal layers of the multilayer crystal layer are each formed of the same scintillation crystal, and each type of scintillation crystal forming each crystal layer is at least two, The stacked structure of the multilayer crystal layer is characterized by the presence of at least one or more offsets between the crystal layers, thus providing more detailed information on the depth of interaction (DOI) between γ-rays and the scintillation crystal compared to conventional scintillation detectors. , Thus more precisely determining the initial location of the γ-rays, and maintaining improved sensitivity while maintaining sensitivity. It is an object of the present invention to provide a scintillation detector having a resolution and a positron emission tomography apparatus using the same.

Positron Emission Tomography, Scintillation Crystal, Scintillator, Photomultiplier, Interaction Depth, DOL, BGO, LSO, LuAP, LuYAP, LYSO, GSO, PET, PMT

Description

Positron emission tomography devices and scintillator detectors for the same

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to imaging devices widely used in medical equipment, and more particularly, to a positron emission tomography (PET) device and a scintillation detector for the tomography device.

PET is computed tomography, such as X-ray computed tomography (CT) and single photon emission computerized tomography (SPECT), which typically emits positrons in vivo for research and diagnosis. The radioactive sample is injected by intravenous injection or inhalation, and then detected. For example, based on the fact that some cancer cells accumulate more glucose than others, FDG, which combines the radioisotope F ¹⁸ with glucose with a half-life of about 110 minutes, is used to track cancer cells. As such, PET has been used for the diagnosis and research of various diseases such as metabolism research of the human body, cancer diagnosis, heart and nervous system abnormalities.

Positron emitting nuclides are predominantly unstable isotopes with a somewhat larger number of neutrons, and nuclides, such as O ¹⁵ , N ¹³ , C ¹¹ and F ¹⁸ , which are mainly used in PET, are stabilized by emitting positrons. By a phenomenon called "pair annihilation", positrons emitted from positron emitting nuclides in the human body combine with nearby electrons to emit γ-rays. According to the law of conservation of energy and momentum of conservation related to the mass-energy equivalence principle, positrons that have reached a stationary state are transformed into extinction gamma rays with 511 keV energy, which are combined with nearby electrons and emitted in opposite directions.

By detecting and analyzing a pair of γ-rays emitted in opposite directions, the occurrence position of γ-rays can be determined. As a result, the frequency of occurrence of γ-rays, that is, the concentration of the labeled sample, can be determined as a function of spatial position coordinates. If the results are displayed using a display means or the like, the distribution of radionuclides in the human body can be known.

Conventional PET devices based on this principle generally adopt a structure in which a plurality of detectors 300 are arranged radially to form an annular detector ring, and the plurality of detector rings are arranged on an axis and the subject is placed therein. After placing it near the center of, measure the three-dimensional spatial distribution of radionuclides.

The most important factors that determine the performance of PET devices are spatial resolution and sensitivity. In order to achieve improved performance, it is also possible to place smaller sized detectors tightly, but there are certain limitations due to the miniaturization of components. Furthermore, the cost increases due to the increase in the number of detectors and electronic readers. have.

The detector for detecting the incidence direction and position of the γ-ray in a PET device can be divided into a semiconductor detector and a scintillation detector according to the detection means, and the present invention relates to a scintillation detector as shown in FIG. Scintillator is a substance that emits light due to collision of radiation, NaI, Bismuth Germanate Oxide (BGO), Lutetium Oxyorthosilicate (LSO), Lutetium-Yttrium Aluminum Perovskite (Luyap) Scintillation crystals such as T) are used in scintillation detectors.

The γ-rays generated at the positions of the radionuclides near the center of the detector ring inner space reach the two scintillation detectors 300 located in opposite directions and form the front crystal layers. Flash signals of different wavelengths are generated depending on the type. The generated scintillation signal is detected by the sensing cells of a photo-multiplier tube (PMT) at the rear end of the detector and the photoelectrons are emitted. The photoelectrons are converted into digital signals that are easily processed and output. The signals are analyzed and reconstructed by a signal processing unit composed of a computer or the like and displayed as a tomographic image containing three-dimensional information through a display device.

2 and 3 are front and right side views showing a conventional laminated structure for improving spatial resolution in two-layer scintillation detectors for PET devices.

The scintillation detector of FIG. 2 is a two-layer scintillation detector 100 in which two crystal layers are formed using two kinds of crystals. When γ-rays are incident on the first crystal layer 121 and the second crystal layer 122 stacked on the plane where the sensing cell 111 of the photomultiplier tube 110 is located, different temporal characteristics are shown in FIG. 6. Based on emitting a scintillation signal of magnitude, one can identify the layer on which the interaction with the γ-rays occurred. By detecting the depth of interaction (DOI) in more detail in comparison with a conventional detector, more detailed measurement and analysis of the direction of incidence of the γ-ray is possible. Therefore, the initial generation position of the γ-line can be determined more precisely, and as a result, the spatial resolution is improved in the analysis of the output signal.

The scintillation detector of FIG. 3 is a two-layer scintillation scintillator 200 in which two crystal layers are formed using the same kind of crystals. By adopting a stacked structure in which a constant offset exists between the first crystal layer 221 and the second crystal layer 222 stacked on the plane where the sensing cell 211 of the photomultiplier tube 210 is located, interaction occurs. Differentiating layers will provide more detailed DOI information.

In addition, various approaches have been made to improve the spatial resolution and sensitivity of PET devices, and despite the gradual improvement, the performance of PET devices remains unsatisfactory. There is a continuing need for improved sensitivity. Moreover, in addition to this limitation of spatial resolution, the larger the "radial extensional distortion", or radial length (vertical distance from the central axis in annular detector rings), which is well known in PET devices, the more spatial resolution at that location becomes. There is a sharply decreasing geometric limit.

In order to solve the problems of the prior art, the multilayer scintillation detector according to the present invention forms three or more crystal layers using different kinds of scintillation crystals, and at the same time, provides a constant offset between the crystal layers of the scintillation detector. By forming the crystal layer, it is possible to detect a flash signal having more detailed DOI information as compared with a conventional flash detector. In particular, the present invention eliminates the disadvantage of having to sacrifice sensitivity in order to improve the spatial resolution when the conventional flash detectors are simply joined, thereby improving the spatial resolution without degrading the sensitivity, especially the radial spatial resolution is significantly improved. An object of the present invention is to provide a detector and a PET device using the same.

In order to achieve the above object, the present invention is characterized in that three or more crystal layers are formed by using two or more kinds of flash crystals in a scintillation detector, and at least one offset is provided between the crystal layer and the crystal layers. do.

As described above, the present invention has the following characteristics and advantages by adopting a laminated structure in which two or more kinds of scintillation crystals are stacked to form a multilayered crystal layer and offset exists between the crystal layers. .

In the positron emission tomography apparatus using the three-layer scintillation detector 300 according to the present invention, the spatial resolution is remarkably improved as compared with the case of using the conventional two-layer scintillation detector 100 as the radius length increases. The performance of the scintillation detector, and also the performance of the positron emission tomography apparatus, is mainly determined by the spatial resolution and sensitivity, and the result shows that the spatial resolution is significantly improved without sacrificing the sensitivity.

The above effect of the three-layer scintillation detector according to the present invention can be explained as it is cited as the effect of the positron emission tomography apparatus including such scintillation detector.

Although the present invention has been described in connection with the above-mentioned preferred embodiments, these embodiments are intended to more clearly disclose the present invention, and various modifications without adding new ones in the scope of the matter described in the specification and drawings. Or it will be possible to make modifications. Accordingly, the appended claims should be understood to include such modifications and variations as may be inferred by those skilled in the art based on the matters described in this specification and drawings and their descriptions.

Hereinafter, with reference to the accompanying drawings showing a preferred embodiment of the scintillation detector according to the present invention will be described for the configuration and operation. 2 and 3 is a laminated structure of a conventional two-layer flash detector, Figure 4 is a laminated structure of a three-layer flash detector according to the present invention, and Figure 5 is a four-layer flash of another embodiment according to the present invention Front and right side views showing the stacked structure of the detector.

As shown in FIG. 4, the three-layer scintillation detector 300 according to the preferred embodiment of the present invention forms a first crystal layer 321 on the plane where the sensing cells 311 of the photomultiplier tube 310 are located. The second crystal layer 322 is formed on the first crystal layer 321 so as to coincide with the first crystal layer 321 without offset, and then a constant offset is placed on the second crystal layer 322 to provide a third crystal layer ( 323 has a laminated structure formed.

The second crystal layer 322 and the third crystal layer 323 are formed of the same kind of flash crystal and the first crystal layer 321 is formed of a different flash crystal. The scintillation crystals forming the first crystal layer 321 and the scintillation crystals constituting the second crystal layer 322 and the third crystal layer 323 include NaI, Bismuth Germanate Oxide (BGO), Lutetium Oxyorthosilicate (LSO), and Lutetium Oxygen. Orthosilicate), LuYAP (Lutetium-Yttrium Aluminum Perovskite, Lutetium-Yttrium Aluminum Perovskite) and the like. As shown in FIG. 6, the flash signal of LSO and LuYAP can be seen that there is a significant difference in pulse shape, which is due to the large difference in the decay time constants of the two crystals. For this purpose, it is preferable to form the first to third crystal layers using scintillation crystals of LSO and LuYAP.

In addition, it is preferable that the total thickness of the stacked three crystal layers is not longer than the thickness of the crystal layers of the conventional two-layer scintillation detector (100,200). This is because the sensitivity is lowered by the increase in the thickness of the crystal layer, and when the thickness of the crystal layers are the same, there is an advantage that the existing PET device can be used as it is by replacing only the crystal layer.

As shown in FIG. 4, the stacked structure of the three-layer scintillation detector 300 according to the present embodiment has N ² rectangular parallelepiped crystals on the plane where the sensing cell 311 of the photomultiplier tube 310 is located. The first crystal layer 321 is formed using N, and the second crystal layer 322 is formed using N ² crystals to coincide with the first crystal layer 321 thereon. Subsequently the second crystal layer 322, the crystal number of the third crystal layer 323 is formed on the (N-1) is ^2, the center of the square forms are second crystal layer at 322, four crystal and the Tricrystalline layer 323 A third crystal layer is formed so that the centers of crystals coincide. Therefore, the third crystal layer is laminated with an offset equal to half the length of the sum of the length of one side of the crystal and the length of the gap between the crystals with respect to the second crystal layer. The sensing cells 311 of the photomultiplier tube 310 may coincide with the first crystal layer 321 and the second crystal layer 322, or may coincide with the third crystal layer 323.

5 is a front view and a right side view showing a four-layer flash detector 400 according to another embodiment of the present invention. Similar to the three-stage scintillation detector 300 of FIG. 4, the four-stage scintillation detector 400 according to the present embodiment uses two or more different kinds of scintillation crystals, and has a common technical characteristic that offsets exist between the crystal layers. Have The stacked structure of the four-layer flash detector 400 is formed so that the first crystal layer 421 and the second crystal layer 422 coincide without offset, and the third crystal layer 423 is the same as described with reference to FIG. 4. The fourth crystal layer 424 is formed to be equal to the third crystal layer 423 without being offset. The scintillation crystal used is preferably the first crystalline layer 421 and the third crystalline layer 423 is LuYAP and the second crystalline layer 422 and the fourth crystalline layer 424 are LSO. The four-layer scintillation detector 400 can achieve improved spatial resolution by providing more detailed DOI information than the three-stage scintillation detector 300, but it involves an increase in complexity and cost in the manufacture of the detector and analysis of output data. do.

Hereinafter, as another aspect of the present invention, a photoelectron emission tomography apparatus using the three-layer scintillation detector 300 will be described. A typical PET device includes a detector including a detector ring, a signal processor, and a display.

The detection unit is formed by forming a plurality of three-layer flash detector 300 described above radially to form a detector ring and stacking one or more detector rings in the axial direction of the detector ring. The subject is mounted in a cylindrical inner space formed by the tip of the third crystal layer 323 of the scintillation detector 300. In particular, the PET device having a radius of 100 to 200 mm in the inner space is widely used for research and diagnosis of gene expression and protein interaction in vivo using small animals.

The signal processing unit includes a computer or other computing device in which a program for analyzing the data output from the detection unit and determining a generation position of the γ-line is embedded. The program used in the present apparatus performs the calculation of determining the DOI of each output pulse based on the waveform of the pulse since the data output from the detector includes DOI information based on the type of crystal and the offset between the crystal layers. The signal processing section and the display section are by means commonly used in conventional positron emission tomography apparatus.

In the plane perpendicular to the axis of the annular detector ring, the radial spatial resolution and the tangential spatial resolution are measured while increasing the radial length by 20 mm, and in the measurement on the small animal positron emission tomography apparatus, In order to compare the spatial resolution of the three-stage scintillation detector 300 and the spatial resolution of the two-stage scintillation detector 100 of FIG. 2 based on the same sensitivity, the total thicknesses of the crystal layers of the scintillation detectors coincide with each other. Do it.

Each of the crystals used for the lamination is a rectangular parallelepiped with a square cross section and a length of 2 mm in length and width. The first crystalline layer 121 and the second crystalline layer 122 of the conventional two-layer scintillation detector 100 are LuYAP and LSO, respectively, and the thickness of the crystal layer (ie, the length of the cuboid crystal) is 9.5 mm. In the three-layer scintillation detector 300 according to the present invention, the first crystal layer 321 is LuYAP, the second crystal layer 322 and the third crystal layer 323 are LSO, and the first crystal layer 321 and The thickness of the second crystal layer 322 is 7 mm and the thickness of the third crystal layer 323 is 5 mm. Therefore, the total thickness of the stacked layers is equal to 19 mm. Each of the crystals in the same crystal layer are all spaced 0.3 mm apart.

The scintillation detectors thus formed are combined to form a detector ring as in a conventional positron emission tomography apparatus, and the four detector rings are arranged in an axial direction to form a detector having a cylindrical inner space in which the subject is placed. It was.

1 is a perspective view showing a schematic form of a scintillation detector.

2 is a front view and a right side view showing a conventional laminated structure of the scintillation detector crystal layer of FIG.

3 is a front view and a right side view showing another conventional laminated structure of the scintillation detector crystal layer;

Figure 4 is a front view and a right side view showing the laminated structure of the three-layer flash detector crystal layer according to an embodiment of the present invention.

5 is a front view and a right side view showing a laminated structure of a four-layer scintillation detector crystal layer according to another embodiment of the present invention.

6 is a graph showing the form of a flash signal of a crystal.

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Explanation of symbols on the main parts of the drawings

100, 200, 300, 400: scintillation detector

110, 210, 310, 410: photomultiplier

111, 211, 311, 411: sensing cell

121, 221, 321, and 421: first crystal layer

122, 222, 322, and 422: second crystal layer

323, 423: third crystal layer

424: fourth crystal layer

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L ₁₁ : LSO flash signal

L ₁₂ : LuYAP80 flash signal

L ₁₃ : LuYAP70 flash signal

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Claims (8)

A photomultiplier tube 310 including sensing cells 311 sensing the scintillation signals from the crystal layers and amplifying and outputting photoelectrons emitted from the sensing cells 311; And And a multi-layered crystal layer having at least three crystal layers sequentially formed on the plane of the photomultiplier tube 310 in which the sensing cells 311 are positioned. The same flash crystal is formed in each of the crystal layers, At least two kinds of scintillation crystals constituting the multilayer crystal layer are used for each layer, In the multilayer crystal layer, the sequentially formed crystal layers are the first crystal layer 321, the second crystal layer 322, and the third crystal layer 323, and the offsets are the second crystal layer 322 and the first crystal layer. It exists between the three crystal layer 323, the first crystal layer 321 is formed of a first crystal, the second crystal layer 322 and the third crystal layer 323 is formed of a second crystal, The first crystal and the second crystal have different characteristics of the flash signal generated in each crystal so that the crystal layer in which the flash signal is generated can be identified. The first crystal is lutetium-yttrium aluminum perovskite (LuYAP), and the second crystal is lutetium oxyorthosilicate (LSO), scintillation detector for a positron emission tomography apparatus. The method of claim 1, The first crystal layer 321 and the second crystal layer 322 are square crystal layers formed of N ² scintillation crystals each having a rectangular parallelepiped shape, and the third crystal layer 323 has a rectangular parallelepiped shape (N-1). ) scintillation detector for positron emission tomography apparatus characterized in that the ^second decision of the square formed by one scintillation crystal layer. The method according to claim 1 or 2, The first crystalline layer 321 and the second crystalline layer 322 are each 7 mm thick, the third crystalline layer 323 is 5 mm thick, the scintillation crystals forming the crystal layer has a length of the horizontal and vertical Scintillation detector for a positron emission tomography apparatus, characterized in that each 2mm. A photomultiplier tube 310 including sensing cells 311 sensing the scintillation signals from the crystal layers and amplifying and outputting photoelectrons emitted from the sensing cells 311; And And a multi-layered crystal layer having at least three crystal layers sequentially formed on the plane of the photomultiplier tube 310 in which the sensing cells 311 are positioned. The same flash crystal is formed in each of the crystal layers, At least two kinds of scintillation crystals constituting the multilayer crystal layer are used for each layer, In the multilayer crystal layer, the sequentially formed crystal layers are the first crystal layer 421, the second crystal layer 422, the third crystal layer 423, and the fourth crystal layer 424. Is present between the bi-crystal layer 422 and the third crystal layer 423, The first crystalline layer 421 and the third crystalline layer 423 are formed of a first crystal, the second crystalline layer 422 and the fourth crystalline layer 424 are formed of a second crystal, and the The first crystal and the second crystal have different characteristics of the flash signal generated in each crystal so that the crystal layer in which the flash signal is generated can be identified. The first crystal is lutetium-yttrium aluminum perovskite (LuYAP), and the second crystal is lutetium oxyorthosilicate (LSO), scintillation detector for a positron emission tomography apparatus. The method of claim 4, wherein The first crystal layer 421 and the second crystal layer 422 are square crystal layers formed of N ² scintillation crystals each having a rectangular parallelepiped shape, and the third crystal layer 423 and the fourth crystal layer 424. scintillation detector is for positron emission tomography apparatus characterized in that the crystal layer of the square formed by each of the rectangular shape (N-1) ² of scintillation crystals. A detector including a detector ring formed of a plurality of flash detectors 300; A signal processor for analyzing the signals output by the plurality of flash detectors 300 and outputting an image signal; And A positron emission tomography apparatus comprising: a display unit for displaying an image signal output from the signal processor, wherein the scintillation detector 300 forming the detector ring comprises: A photomultiplier tube 310 including sensing cells 311 sensing the scintillation signals from the crystal layers and amplifying and outputting photoelectrons emitted from the sensing cells 311; And And a multi-layered crystal layer having at least three crystal layers sequentially formed on one surface of the photomultiplier tube where the sensing cells 311 are positioned. The same scintillation crystal is formed in each of the crystal layers, and the kind of scintillation crystals constituting the multilayer crystal layer is at least two or more per layer, In the scintillator detector 300, the sequentially formed crystal layers are the first crystal layer 321, the second crystal 322 layer, and the third crystal layer 323. It exists between the three crystal layer 323, the first crystal layer 321 is formed of a first crystal, the second crystal layer 322 and the third crystal layer 323 is formed of a second crystal, The first crystal and the second crystal have different characteristics of the flash signal generated in each crystal so that the crystal layer in which the flash signal is generated can be identified. And the first crystal is lutetium-yttrium aluminum perovskite (LuYAP) and the second crystal is lutetium oxyorthosilicate (LSO). The method of claim 6, The scintillation detector 300 of the first crystal layer 321 and the second crystal layer 322 and the crystal layer of the square is formed in each rectangular shape in N ² of scintillation crystals, and the third crystal layer 323 positron emission tomography apparatus characterized in that the rectangular shape of (N-1) the ^second decision of the square formed by one scintillation crystal layer. The method of claim 7, wherein Positron emission tomography apparatus, characterized in that the inner space of the detector ring formed by the tip of the third crystal layer (323) of the scintillation detector (300) has a radius of 50mm to 500mm.

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