CN110974267B - PET detector of composite crystal array and construction method thereof - Google Patents
PET detector of composite crystal array and construction method thereof Download PDFInfo
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Abstract
The invention discloses a PET detector of a composite crystal array and a construction method thereof, wherein the structure of the PET detector of the composite crystal array comprises the following components: the device comprises a scintillation crystal module, a photoelectric conversion device module and a detector electronics module, wherein the scintillation crystal module is used for converting gamma photons into visible light photons and soft ultraviolet light photons and transmitting the visible light photons and the soft ultraviolet light photons to the photoelectric conversion device module, and the scintillation crystal module consists of an A crystal module and a B crystal module to form an AB type mixed light and heavy rare earth single crystal array; the A crystal module is light rare earth single crystal, and the B crystal module is heavy rare earth single crystal; and a and B form a crystal array using a configuration of a partition structure. The scintillation crystal array in the PET detector of the composite crystal array consists of two mixed light and heavy rare earth crystals AB, so that the lutetium content in the PET detector is reduced, and the cost and the system complexity of the PET detector are reduced on the premise of ensuring the high performance of the PET detector.
Description
Technical Field
The invention relates to the technical field of electronic information and medical instruments, in particular to a PET detector of a composite crystal array and a construction method thereof.
Background
Positron Emission Tomography (PET) is a non-invasive in vivo imaging method, which can non-invasively, quantitatively and dynamically evaluate the metabolic level, biochemical reaction, functional activity and perfusion of various organs in the human body, represents the advanced technology in the field of current medical imaging, and is a clinical functional imaging device with the highest sensitivity. The detector is the main part of the whole PET, and consists of a scintillation crystal, a photoelectric conversion device and reading electronics, wherein the performance of the scintillation crystal determines the performance of the detector, and the scintillation crystal detector is the core of the system. The ideal crystal material should have the characteristics of sufficiently high density, short afterglow time, high light output, good energy resolution, low production cost and the like. The high density and high atomic number can effectively improve the detection efficiency of gamma rays; the short afterglow time can better perfect time matching and reduce random counting; the high light output can increase the number of crystals per photodetector; the good energy resolution can reduce image scattering and make the image clearer. These properties are in turn mutually limiting with respect to the detector crystal. The scintillation crystal is combined with a photoelectric detector, and can be widely applied to the detection of high-energy photons and particles, such as high-energy physics, well logging, safety, medical imaging and the like. Since the birth of inorganic scintillation crystals has been the history of nearly 20 years till now, hundreds of crystals are researched in sequence, the emergence of novel scintillation crystals does not mean the disappearance of the traditional scintillation crystals, the growth process and the quality of the traditional crystals are relatively stable, the application in the market is approved, and therefore the performance needs to be further improved, and the new crystals need to be developed and are equally important. From the viewpoint of the detection of reserves and the delivery price, various rare earth elements are extremely unbalanced, so the balanced utilization of the rare earth becomes a problem to be solved urgently. The PET of high-end medical equipment has great use amount of heavy rare earth element lutetium, and the lutetium is low in reserve and high in price, so that the cost of high-performance PET is difficult to reduce.
Accordingly, there is a need for an improved PET detector that addresses the above-mentioned problems with crystals to overcome the noted deficiencies in the prior art.
Disclosure of Invention
In order to overcome the defects of the prior art, the invention aims to provide a PET detector of a composite crystal array and a construction method thereof, solve the problems of the prior art, reduce the lutetium content in the PET detector and reduce the cost and the system complexity of the PET detector on the premise of ensuring the high performance of the PET detector.
In order to achieve the purpose, the invention adopts the following technical scheme:
the invention provides a PET detector of a composite crystal array, which is characterized in that the structure of the PET detector of the composite crystal array comprises: the device comprises a scintillation crystal module, a photoelectric conversion device module and a detector electronics module, wherein the scintillation crystal module consists of an A crystal module and a B crystal module to form an AB type mixed light and heavy rare earth single crystal array; the A crystal module is light rare earth single crystal, and the B crystal module is heavy rare earth single crystal; and a and B form a crystal array using a configuration of a partition structure.
As a further improvement of the scheme, the A crystal element is any one of lanthanum La, cerium Ce, praseodymium Pr, neodymium Nd, promethium Pm, samarium Sm, europium Eu and gadolinium Gd8 in lanthanide; the B crystal element is any one of 9 elements of terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu, scandium Sc and yttrium Y.
As a further improvement of the scheme, in the crystal array, the partition structure adopts a surrounding partition structure, 1B crystal is surrounded by 8A crystals to form a 3X 3 Sudoku crystal array, wherein the A crystals and the B crystals have the same shape and size and are both square crystals.
As a further improvement of the scheme, the number of the crystals is increased by analogy with the nine-grid type crystal array, so that an n multiplied by n crystal array is formed, and the n multiplied by n crystal array is formed by taking a 3 multiplied by 3 nine-grid type crystal array as a basic unit.
As a further improvement of the scheme, in the crystal array, the partition structure adopts an embedded partition, and a crystal B is placed between the crystal A and the crystal A to form an embedded crystal array, wherein the crystal A and the crystal B have different shapes and sizes.
The embedded crystal array is formed by embedding a layer of B crystals in any two adjacent A crystals by adopting embedded partition, and the number of the crystals is increased by the same method to form an n multiplied by n crystal array, wherein the shapes and the sizes of the A crystals and the B crystals are different.
As a further improvement of the scheme, the corners of the crystal module A and the crystal module B are both right angles; the detector formed by the crystal module A and the crystal module B is packaged into a rectangular detector; when the PET detector of the composite crystal array is packaged, the upper, lower, left, right, front five surfaces are wrapped by reflecting materials.
As a further improvement of the scheme, 8 packaged PET detectors with self-locking structures are respectively fixed on the detector ring, and the detector module is enabled to be provided with 8 limiting structures which are convenient to insert and withdraw along the radial direction.
As a further improvement of the scheme, the embedded scintillation crystal array is formed by combining any one of cerium-doped light rare earth single crystals and any one of heavy rare earth single crystals under the condition of meeting chemical conditions.
As a further improvement of the scheme, the photoelectric conversion device module is used for a device which is responsible for converting optical signals into analog electric signals in the PET detector and is composed of a photomultiplier tube, and the photomultiplier tube is sensitive to the emission spectrum of the scintillation crystal and provides proper signal amplification factor and excellent time response speed; the detector electronics module is responsible for extracting event information from the simulated electrical pulse signals in the PET detector, and the detector electronics can extract information (time, energy, location) of single pulse events and classify them as paired coincidence events according to the single pulse information.
A method for constructing the PET detector of the composite crystal array comprises the following signal transmission steps:
step S1: gamma photons are emitted into the composite crystal array, and the crystal array absorbs energy to deposit and convert the energy into visible light;
step S2: visible light enters the photoelectric conversion module and is converted into an analog electric signal by the photomultiplier;
step S3: the analog electric signal enters a detector electronics module, and the detector electronics converts the analog electric signal into a digital electric signal and obtains time information and energy information in the digital electric signal;
step S4: and carrying out data processing and image reconstruction on the obtained time information and energy information by using computer software to finally obtain a reconstructed image.
The invention has the beneficial effects that:
the scintillation crystal array in the PET detector of the composite crystal array consists of two mixed light and heavy rare earth crystals AB, so that the lutetium content in the PET detector is reduced, and the cost and the system complexity of the PET detector are reduced on the premise of ensuring the high performance of the PET detector. Meanwhile, because the density of the scintillation crystals of two different materials is different, the blocking capability of the solid scintillator with high density to the high-energy particles is stronger, and the blocking capability of the liquid scintillator with low density to the high-energy particles is weaker, so that the scintillator in the radiation sensing device with the two-state structure has the collimation effect on different high-energy particles, and the anti-scattering effect is achieved.
The invention not only reduces the system cost, but also improves the time resolution and the space resolution, and provides a new idea for the balanced utilization of the rare earth. Moreover, the cost of the PET detector is greatly dependent on the cost of the internal scintillation crystal, so the invention also plays a role in promoting the development and the progress of high-end medical equipment such as the PET detector.
The crystal array provided by the two schemes shows a simple but inconstant sense of condensation, and has the characteristics of reasonable structure, beautiful appearance, scientific and reasonable structure and the like, and has higher requirement on the structural precision than the traditional crystal array.
Drawings
Fig. 1 is a schematic structural diagram of an embedded scintillation crystal array of a PET detector of a composite crystal array according to an embodiment of the present invention.
FIG. 2 is a structural diagram of a surrounding scintillation crystal array of a PET detector of a composite crystal array according to an embodiment of the present invention.
Fig. 3 is a schematic structural diagram of an AB crystal module of an embedded scintillation crystal of the present invention.
FIG. 4 is a schematic diagram of a structure of a surrounding type scintillation crystal AB crystal module of the invention.
FIG. 5 is a package structure diagram of an embedded scintillation crystal array of the present invention.
FIG. 6 is a package structure diagram of the surrounding scintillation crystal array of the present invention.
FIG. 7 is a schematic of the geometry of a continuous crystal module in the detector of the present invention.
Fig. 8 is an overall configuration diagram of the detector of the present invention.
Fig. 9 is a schematic diagram of signal transmission in the detector of the present invention.
In the figure:
1. embedded crystal array 2, surrounding crystal array 3, light rare earth single crystal A4, heavy rare earth single crystal B7, embedded crystal array package 8, surrounding crystal array package 9, detector ring 10, photomultiplier 11, detector electronics 12, image reconstruction and imaging module
Detailed Description
The technical scheme of the invention is further explained by the specific implementation mode in combination with the attached drawings. It should be understood that in the description of the present application, the terms "center", "length", "depth", "thickness", "upper", "lower", "front", "rear", "left", "right", "vertical", "horizontal", "top", "bottom", "inner", "outer", "clockwise", "counterclockwise", etc. indicate orientations and positional relationships based on those shown in the drawings, and are only for convenience of description and simplicity of description, but do not indicate or imply that the referenced device or element must have a particular orientation, be constructed and operated in a particular orientation, and therefore should not be considered as limiting the present application.
The invention provides a PET detector of a composite crystal array, which structurally comprises: the device comprises a scintillation crystal module, a photoelectric conversion device module and a detector electronics 11 module, wherein the scintillation crystal module is used for converting gamma photons into visible light photons and soft ultraviolet photons and transmitting the visible light photons and the soft ultraviolet photons to the photoelectric conversion device module, and the scintillation crystal module is an AB type mixed light and heavy rare earth single crystal array consisting of an A crystal module and a B crystal module; the A crystal module is light rare earth single crystal, and the B crystal module is heavy rare earth single crystal; and a and B form a crystal array using a configuration of a partition structure. The A crystal module is an A light rare earth single crystal 3, and the B crystal module is a B heavy rare earth single crystal 4; the A crystal element is any one of La, Ce, Pr, Nd, Pm, Sm, Eu, Gd8 elements in lanthanide; the B crystal element is any one of 9 elements of terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu, scandium Sc and yttrium Y.
The rare earth crystal is a crystal in which rare earth elements can completely occupy a certain lattice point in a crystallographic structure, and is widely applied to laser technology and ionizing radiation detection technology as a core working substance. The mixed light and heavy rare earth single crystal is characterized in that only heavy rare earth elements are used in a crystal array instead of the traditional sense, light and heavy rare earth elements are mixed, wherein light rare earth elements refer to lanthanum La, cerium Ce, praseodymium Pr, neodymium Nd, promethium Pm, samarium Sm, europium Eu and gadolinium Gd8 elements in lanthanide elements, and heavy rare earth elements refer to terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu7 elements in lanthanide elements, and two elements scandium Sc and yttrium Y closely related to the lanthanide elements.
The surfaces of the AB two crystal modules have different treatment modes. Because after gamma photons are deposited in the crystal to generate scintillation photons, the scintillation photons can propagate in the crystal and are reflected, refracted or lost on the surface of the crystal, a part of the photons are finally detected by the SiPM detection unit array coupled to the bottom surface of the crystal, and the positioning of the action positions of the gamma photons and the crystal is mainly realized through the distribution of output signals of the bottom SiPM detection array. The crystal surface processing mode directly determines the behavior of scintillation photons on the crystal surface and finally influences the distribution of output signals of the SiPM detection array, so the crystal surface processing mode has a crucial influence on the three-dimensional position positioning performance of the detector module. Different scintillation photon transport processes may also result in different light collection efficiencies of the detector modules, and thus the crystal surface treatment may also have an impact on the temporal resolution and energy resolution performance of the detector modules. In summary, a suitable crystal surface treatment may play an important role in improving the performance of the continuous crystal detector module. The crystal surface treatment mainly comprises how to treat the crystal surface and what kind of reflective film is selected to wrap the crystal. The conventional crystal surface treatment mainly includes polishing and carborundum grinding, and the conventional crystal reflective film mainly includes ESR reflective film, Teflon reflective film and the like. The invention not only reduces the system cost, but also improves the time resolution and the space resolution, and provides a new idea for the balanced utilization of the rare earth.
The internal basic principle of the PET detector of the composite crystal array is as follows: the radionuclide decays to generate positrons, which combine with negative electrons to generate annihilation radiation, and two gamma photons with 511keV and opposite flight directions are generated. These two photons have very important properties: the simultaneous temporal and almost opposite flying-off occurs, which makes it possible to use two detectors placed opposite each other outside the body, detecting them with coincidence techniques, which are now commonly measured by closed multiloop detectors. If within a specified time window (typically 0 ns-15 ns) the detector detects two photons at 180 degrees (0.25 degrees) to each other, i.e. a coincidence event, the annihilation point is on the line between the two crystal blocks where the flash occurred, forming a lor (line Of reaction) response line, which is recorded in memory in coincidence. Since the paths of the two photons in the body are different, the times of arrival at the respective detectors are also different, and the position of the annihilation point on the coincidence line can be calculated. The coincidence circuit system performs coincidence measurement to generate original data, and the computer system is used for completing data acquisition, system monitoring, attenuation correction and image reconstruction.
According to one embodiment of the invention, as shown in fig. 1, a first embodiment of a method for constructing a composite crystal array PET detector and a method for constructing the same, namely an embedded scintillation crystal array, comprises an a crystal and a B crystal. The crystal A is a crystal made of cerium-doped light rare earth single-crystal lanthanum bromide, and the crystal B is a crystal made of heavy rare earth single-crystal lutetium yttrium silicate. A layer of B crystals are embedded in any two adjacent A crystals, and the number of the crystals is increased by the same method to form an n multiplied by n crystal array which is an embedded scintillation crystal array.
According to an embodiment of the invention, as shown in fig. 2, a PET detector of a composite crystal array and a second construction method of the construction method, namely a surrounding crystal array 2, comprises an a crystal and a B crystal. The crystal A is a crystal made of cerium-doped light rare earth single-crystal lanthanum bromide, and the crystal B is a crystal made of heavy rare earth single-crystal lutetium yttrium silicate. Eight A crystals surround a B crystal in the middle to form a 3 x 3 nine-grid type crystal array, and the number of the nine-grid type crystals is increased by the same method to form an n x n crystal array which is a surrounding type scintillation crystal array.
According to one embodiment of the present invention, as shown in FIG. 3, the cerium-doped light rare earth single crystal lanthanum bromide A crystal has a length and width of 1mm and a height of 3mm, and the embedded heavy rare earth single crystal yttrium lutetium silicate B crystal has a length of 1mm, a width of 0.2mm and a height of 3 mm.
According to one embodiment of the present invention, as shown in FIG. 4, the cerium-doped light rare earth single crystal lanthanum bromide A crystal has a length and width of 1mm and a height of 3mm, and the embedded heavy rare earth single crystal yttrium lutetium silicate B crystal has a length and width of 1mm and a height of 3 mm.
According to one embodiment of the invention, as shown in fig. 5, the embedded scintillation crystal array is a 6 x 6 array with a layer of encapsulation on the outer surface of the array, and the embedded crystal array encapsulation 7 is composed of a reflective film, a coupling layer and SiPM detection units. The packaged shape is cubic.
According to one embodiment of the invention, as shown in fig. 6, the surrounding scintillation crystal array is a 6 x 6 array with a layer of encapsulation on the outer surface of the array, and the surrounding crystal array encapsulation 8 is composed of a reflective film, a coupling layer and SiPM detection units. The packaged shape is cubic.
According to an embodiment of the present invention, the crystal surface in fig. 7 is covered with a reflective film, the reflective film is made of teflon and aluminum foil for preventing the scintillation photons from refraction and reflection in the scintillation crystal, there is a coupling layer close to the inside of the reflective film, the coupling layer functions like a light guide to prevent the photons from refracting too much in the air to enter the photomultiplier 10, there is a SiPM detection unit at the bottom of the crystal, which is the front end of the photomultiplier 10.
When the PET detector of the composite crystal array is packaged, the upper, lower, left, right, front five surfaces are wrapped by reflecting materials such as Teflon, BaSO4 and the like on the surfaces, so that the purpose of doing so is that scintillation photons can be refracted and reflected in the scintillation crystal, and certain photons can be lost after leaving the scintillator and entering the light guide or the silicone layer. The difference in optical path length has an effect on the temporal characteristics of the scintillation pulse. The reduction in photon counts during transport affects the energy characteristics of the scintillation pulse. In particular, the way the crystal is processed can have a significant effect on the reflection and refraction of scintillation photons. If the surface of the crystal is wrapped by the reflecting material, a part of the emitted photons can be reflected back to the crystal and continue to be transmitted in the crystal.
According to an embodiment of the present invention, as shown in fig. 8, the PET detector of the composite crystal array disclosed in this embodiment includes: a scintillation crystal module comprising a plurality of crystal modules, a photomultiplier module, and a detector electronics module, wherein the detector electronics module comprises first module electronics and second module electronics.
Furthermore, the scintillation crystal module comprises an A crystal module and a B crystal module, and both the A crystal module and the B crystal module are used for absorbing gamma photons emitted in a living body, converting the gamma photons into visible light photons and soft ultraviolet photons and transmitting the visible light photons and the soft ultraviolet photons to a photocathode module in the photomultiplier module;
the photomultiplier module comprises a photocathode module, a focusing electrode module, a dynode module and a photoanode module, wherein the photocathode module is used for converting photons transmitted by the crystal module A and the crystal module B into photoelectrons through a photoelectric effect and transmitting the photoelectrons to the focusing electrode module; the focusing electrode module is used for focusing the photoelectron beam and then transmitting the focused photoelectron beam to the dynode module, the dynode module carries out secondary emission on the photoelectron beam and multiplies the photoelectron beam to the photoanode module, and the photoanode module collects amplified electrons by using an anode and outputs the collected electrons as signals to the reading electronic module. Furthermore, the reading electronics module comprises a first electronics module and a second electronics module, the first electronics module comprises an amplifier module and an analog-to-digital conversion module, and the second electronics module comprises an MVT module, an FPGA module and an Ethernet module. The amplifier module and the MVT module receive electric signals from the photoelectric anode module, then the amplifier module amplifies the electric signals and transmits the electric signals to the analog-to-digital conversion module, the analog-to-digital conversion module performs analog-to-digital conversion on the amplified analog electric signals to obtain digital electric signals and transmits the digital electric signals to the FPGA module, the MVT module sparsely quantizes the electric signals and transmits the electric signals to the FPGA module, the FPGA module processes the signals and transmits the processed electric signals to the Ethernet module, and the Ethernet module processes the signals and transmits the processed electric signals to the data preprocessing module in the image reconstruction and imaging module; the image reconstruction and imaging module 12 includes a data preprocessing module, an image reconstruction module, and an image post-processing and display module, wherein the digital signal obtained from the ethernet module is transmitted to the data preprocessing module, the data preprocessing module preprocesses the signal and transmits the signal to the image reconstruction module, and the image reconstruction module transmits the processed signal to the image post-processing and display module to finally obtain a reconstructed image.
As shown in fig. 9, a single crystal array is placed in a pre-designed spacing structure in the detector ring 9, the size of the spacing structure being equal to that of a single detector cube module. The detector is the main part of the whole positron emission imaging system and consists of a scintillation crystal, a photomultiplier tube 10 and detector electronics 11. The scintillation crystal and the photomultiplier tube 10 are fixed on the detector ring 9, the crystal on the detector ring 9 converts high-energy photons into visible light, the photomultiplier tube 10 connected with the scintillation crystal converts optical signals into electrical signals, the electrical signals are converted into time pulse signals, the time coupling of each crystal pulse signal is checked and judged according to a circuit, the interference of rays from other sources is eliminated, and the position of a positron is given out through operation. Each crystal on a closed ring detector has a coincidence relationship with an opposing set of crystals to form a set of coincidence lines of fan-shaped beams, the intersection of all fan-shaped beams determining the radial field of view (FOV) of the detector, and the method of achieving collimation without a shielded collimator and relying instead on a special direction of the two photons and coincidence circuits, called "electronic collimation", whenever two photons generated by annihilation radiation occurring within this field of view cannot strike the same crystal "simultaneously" (within the same time window). PET employs coincidence detection techniques for electronic collimation, which greatly reduces random coincidence events and background, and the electronic collimator has very high sensitivity (without the effects of lead shielding) and resolution, and only two photons captured simultaneously by crystals 180 ° from each other within a specified time window can become a coincidence event. The coincidence circuit confirms that coincidence events entering the same time window come from one annihilation, the coincidence events are stored according to a specified projection plane, then the computer completes image reconstruction by adopting the technologies of scattering, accidental coincidence signal correction, photon flight time calculation and the like, and the reconstructed image shows the distribution of the tracer in a human body.
The corners of the crystal module A and the crystal module B are right angles; the detector formed by the crystal module A and the crystal module B is packaged into a rectangular detector; when the PET detector of the composite crystal array is packaged, the upper, lower, left, right, front five surfaces are wrapped by reflecting materials. The eight encapsulated PET detectors with self-locking structures are respectively fixed on the detector ring 9 and make the detector modules convenient to be inserted and pulled along the radial direction.
A method for constructing the PET detector of the composite crystal array comprises the following signal transmission steps: step S1: gamma photons are emitted into the composite crystal array, and the crystal array absorbs energy to deposit and convert the energy into visible light; step S2: visible light enters the photoelectric conversion module and is converted into an analog electric signal by the photomultiplier; step S3: the analog electric signal enters a detector electronics module, and the detector electronics converts the analog electric signal into a digital electric signal and obtains time information and energy information in the digital electric signal; step S4: and carrying out data processing and image reconstruction on the obtained time information and energy information by using computer software to finally obtain a reconstructed image.
keV gamma photons generated by positron annihilation strike a scintillation crystal in a scintillation detector, where the gamma photons deposit energy and are then converted into a multitude of relatively low energy scintillation photons. Among them, the scintillation crystal is a functional material that can effectively absorb high-energy rays (X-rays, gamma rays) or high-energy particles and emit ultraviolet and visible light. According to two embodiments of the invention, the crystals used by the AB two crystal modules are yttrium lutetium silicate LYSO and cerium-doped lanthanum bromide LaBr3Both of which are novel crystals, wherein the lutetium yttrium silicate LYSO belongs to the monoclinic system and the space group is C2The melting point is as high as 2050 ℃, the temperature characteristic and the physical and chemical properties are excellent, the resolution is high, and the response can be quicker; bromination of cerium-doped compoundsLanthanum LaBr3Belongs to a hexagonal system, has obvious anisotropy, has the advantages of high light output, fast decay time, high energy resolution, wide application and the like, and is one of the best inorganic scintillators in the current commerce.
The PET detector of the composite crystal array and the construction method thereof are provided aiming at the problems that various rare earth elements are extremely unbalanced, the consumption of heavy rare earth element lutetium in high-end medical equipment PET is extremely large, the lutetium raw material is low in storage amount and high in price, and further the cost of high-performance PET is difficult to reduce. The partition design of the mixed light and heavy rare earth single crystal is feasible, on one hand, the afterglow constants of the light and heavy rare earth single crystals are different, and higher spatial resolution can be obtained by introducing pulse shape information; on the other hand, the light rare earth is used for replacing the heavy rare earth to absorb radiation, so that the system cost can be reduced. Under the verification of PET design simulation software GATE, the spatial resolution of the partition scheme I, namely the embedded crystal array 1 shown in figure 1, is improved by 31.2%, and the cost calculated by the latest rare earth market trading price is reduced by 40.1%. In addition, due to the application of the cerium-doped light rare earth single crystal LaBr3, 29.3 percent of gamma events exist, and the time resolution is optimized by 2.12 times; with 50.2% gamma events, the temporal resolution is optimized by a factor of 1.53. And the spatial resolution of the second partition scheme, namely the surrounding crystal array shown in the figure 2, is improved by 44.2 percent, and the cost calculated by the latest rare earth market trading price is reduced by 10.3 percent. Similar to scenario one, there are 17.3% gamma events with a time resolution optimized by a factor of 2.19, and 31.9% gamma events with a time resolution optimized by a factor of 1.43. As a preliminary exploration, the design method of the detector provides a new idea for balanced utilization of rare earth in PET.
While the invention has been described with reference to a preferred embodiment, it will be understood by those skilled in the art that various changes may be made and equivalents may be substituted for elements thereof without departing from the spirit and scope of the invention. The present invention is not to be limited by the specific embodiments disclosed herein, and other embodiments that fall within the scope of the claims of the present application are intended to be within the scope of the present invention.
Claims (6)
1. A PET detector of a composite crystal array characterized by: the structure of the PET detector of the composite crystal array comprises: the device comprises a scintillation crystal module, a photoelectric conversion device module and a detector electronics module, wherein the scintillation crystal module is used for converting gamma photons into visible light photons and soft ultraviolet light photons and transmitting the visible light photons and the soft ultraviolet light photons to the photoelectric conversion device module, and the scintillation crystal module consists of an A crystal module and a B crystal module to form an AB type mixed light and heavy rare earth single crystal array; the A crystal module is light rare earth single crystal, and the B crystal module is heavy rare earth single crystal; and A and B form a crystal array using a configuration of a partition structure; the partition structure comprises an enclosed partition;
the surrounding partition wall comprises: using eight A crystals to surround a B crystal in the middle to form a 3X 3 nine-grid crystal array, wherein the A crystal and the B crystal have the same shape and size and are square crystals;
the A crystal element is any one of La, Ce, Pr, Nd, Pm, Sm, Eu, Gd8 elements in lanthanide; the B crystal element is any one of 9 elements of terbium Tb, dysprosium Dy, holmium Ho, erbium Er, thulium Tm, ytterbium Yb, lutetium Lu, scandium Sc and yttrium Y.
2. The composite crystal array PET detector of claim 1, wherein: the surrounding partition is analogized by the nine-grid type crystal array, the number of crystals is increased to form an n multiplied by n crystal array, and the n multiplied by n crystal array is formed by taking a 3 multiplied by 3 nine-grid type crystal array as a basic unit.
3. The composite crystal array PET detector of claim 1, wherein: the corners of the crystal module A and the crystal module B are right angles; the detector formed by the crystal module A and the crystal module B is packaged into a rectangular detector; when the PET detector of the composite crystal array is packaged, the upper, lower, left, right, front five surfaces are wrapped by reflecting materials.
4. The composite crystal array PET detector of claim 3, wherein: the 8 packaged PET detectors with self-locking structures are respectively fixed on the detector ring and enable the detector module to be convenient for drawing and inserting along the radial direction, and 8 limiting structures are arranged on the detector ring.
5. The composite crystal array PET detector of claim 1, wherein: the photoelectric conversion device module is used for converting an optical signal into an analog electric signal in the PET detector and is composed of a photomultiplier tube, and the photomultiplier tube is sensitive to the emission spectrum of the scintillation crystal and provides proper signal amplification and excellent time response speed; the detector electronics module is responsible for extracting event information from the simulated electrical pulse signals in the PET detector, and the detector electronics can extract single pulse event information and classify the single pulse event information into paired coincidence events according to the single pulse information.
6. A method of constructing a PET detector of a composite crystal array as claimed in any one of claims 1 to 5, characterized in that: the method comprises the following signal transmission steps:
step S1: gamma photons are emitted into the composite crystal array, and the composite crystal array absorbs energy to deposit and convert the energy into visible light;
step S2: visible light enters the photoelectric conversion module and is converted into an analog electric signal by the photomultiplier;
step S3: the analog electric signal enters a detector electronics module, and the detector electronics converts the analog electric signal into a digital electric signal and obtains time information and energy information in the digital electric signal;
step S4: and carrying out data processing and image reconstruction on the obtained time information and energy information by using computer software to finally obtain a reconstructed image.
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