CN113671553B - X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector - Google Patents
X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector Download PDFInfo
- Publication number
- CN113671553B CN113671553B CN202111030092.4A CN202111030092A CN113671553B CN 113671553 B CN113671553 B CN 113671553B CN 202111030092 A CN202111030092 A CN 202111030092A CN 113671553 B CN113671553 B CN 113671553B
- Authority
- CN
- China
- Prior art keywords
- layer
- scintillator
- pixels
- pixel
- array
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Active
Links
- 238000001228 spectrum Methods 0.000 title claims abstract description 37
- 238000001514 detection method Methods 0.000 title claims abstract description 16
- 238000004519 manufacturing process Methods 0.000 title claims abstract description 11
- 238000012545 processing Methods 0.000 claims abstract description 13
- 238000005516 engineering process Methods 0.000 claims abstract description 12
- 239000010410 layer Substances 0.000 claims description 185
- 235000012431 wafers Nutrition 0.000 claims description 50
- 238000001914 filtration Methods 0.000 claims description 39
- 238000013461 design Methods 0.000 claims description 26
- 239000010409 thin film Substances 0.000 claims description 23
- 230000008878 coupling Effects 0.000 claims description 15
- 238000010168 coupling process Methods 0.000 claims description 15
- 238000005859 coupling reaction Methods 0.000 claims description 15
- 239000000463 material Substances 0.000 claims description 14
- 238000000034 method Methods 0.000 claims description 14
- 239000002355 dual-layer Substances 0.000 claims description 12
- 239000010408 film Substances 0.000 claims description 11
- SBIBMFFZSBJNJF-UHFFFAOYSA-N selenium;zinc Chemical compound [Se]=[Zn] SBIBMFFZSBJNJF-UHFFFAOYSA-N 0.000 claims description 11
- 238000000926 separation method Methods 0.000 claims description 8
- 238000005498 polishing Methods 0.000 claims description 6
- 230000003287 optical effect Effects 0.000 claims description 5
- 239000002994 raw material Substances 0.000 claims description 5
- 238000005520 cutting process Methods 0.000 claims description 4
- 238000003754 machining Methods 0.000 claims description 4
- 239000003292 glue Substances 0.000 claims description 3
- 238000000227 grinding Methods 0.000 claims description 3
- 238000003384 imaging method Methods 0.000 abstract description 6
- 230000010354 integration Effects 0.000 abstract description 4
- MCVAAHQLXUXWLC-UHFFFAOYSA-N [O-2].[O-2].[S-2].[Gd+3].[Gd+3] Chemical compound [O-2].[O-2].[S-2].[Gd+3].[Gd+3] MCVAAHQLXUXWLC-UHFFFAOYSA-N 0.000 description 11
- 238000006243 chemical reaction Methods 0.000 description 11
- 238000010586 diagram Methods 0.000 description 5
- 230000007547 defect Effects 0.000 description 3
- 239000002356 single layer Substances 0.000 description 3
- 239000000758 substrate Substances 0.000 description 3
- VYPSYNLAJGMNEJ-UHFFFAOYSA-N Silicium dioxide Chemical compound O=[Si]=O VYPSYNLAJGMNEJ-UHFFFAOYSA-N 0.000 description 2
- GWEVSGVZZGPLCZ-UHFFFAOYSA-N Titan oxide Chemical compound O=[Ti]=O GWEVSGVZZGPLCZ-UHFFFAOYSA-N 0.000 description 2
- 230000006978 adaptation Effects 0.000 description 2
- 230000009286 beneficial effect Effects 0.000 description 2
- 230000005540 biological transmission Effects 0.000 description 2
- 230000000694 effects Effects 0.000 description 2
- 238000012986 modification Methods 0.000 description 2
- 230000004048 modification Effects 0.000 description 2
- PFNQVRZLDWYSCW-UHFFFAOYSA-N (fluoren-9-ylideneamino) n-naphthalen-1-ylcarbamate Chemical compound C12=CC=CC=C2C2=CC=CC=C2C1=NOC(=O)NC1=CC=CC2=CC=CC=C12 PFNQVRZLDWYSCW-UHFFFAOYSA-N 0.000 description 1
- 238000003491 array Methods 0.000 description 1
- 229910001632 barium fluoride Inorganic materials 0.000 description 1
- 238000003759 clinical diagnosis Methods 0.000 description 1
- 229910052681 coesite Inorganic materials 0.000 description 1
- 229910052906 cristobalite Inorganic materials 0.000 description 1
- 238000011161 development Methods 0.000 description 1
- 238000002059 diagnostic imaging Methods 0.000 description 1
- 238000000295 emission spectrum Methods 0.000 description 1
- 238000002955 isolation Methods 0.000 description 1
- 238000001748 luminescence spectrum Methods 0.000 description 1
- 238000004020 luminiscence type Methods 0.000 description 1
- 238000005457 optimization Methods 0.000 description 1
- 238000011160 research Methods 0.000 description 1
- 239000004065 semiconductor Substances 0.000 description 1
- 230000035945 sensitivity Effects 0.000 description 1
- 239000000377 silicon dioxide Substances 0.000 description 1
- 235000012239 silicon dioxide Nutrition 0.000 description 1
- 230000003595 spectral effect Effects 0.000 description 1
- 229910052682 stishovite Inorganic materials 0.000 description 1
- 239000000126 substance Substances 0.000 description 1
- 229910052905 tridymite Inorganic materials 0.000 description 1
Classifications
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/161—Applications in the field of nuclear medicine, e.g. in vivo counting
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/1603—Measuring radiation intensity with a combination of at least two different types of detector
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/2002—Optical details, e.g. reflecting or diffusing layers
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/2008—Measuring radiation intensity with scintillation detectors using a combination of different types of scintillation detectors, e.g. phoswich
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/20—Measuring radiation intensity with scintillation detectors
- G01T1/202—Measuring radiation intensity with scintillation detectors the detector being a crystal
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/24—Measuring radiation intensity with semiconductor detectors
- G01T1/243—Modular detectors, e.g. arrays formed from self contained units
-
- G—PHYSICS
- G01—MEASURING; TESTING
- G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
- G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
- G01T1/16—Measuring radiation intensity
- G01T1/24—Measuring radiation intensity with semiconductor detectors
- G01T1/246—Measuring radiation intensity with semiconductor detectors utilizing latent read-out, e.g. charge stored and read-out later
Landscapes
- Physics & Mathematics (AREA)
- Health & Medical Sciences (AREA)
- Life Sciences & Earth Sciences (AREA)
- General Physics & Mathematics (AREA)
- High Energy & Nuclear Physics (AREA)
- Molecular Biology (AREA)
- Spectroscopy & Molecular Physics (AREA)
- Engineering & Computer Science (AREA)
- Biomedical Technology (AREA)
- General Health & Medical Sciences (AREA)
- Medical Informatics (AREA)
- Nuclear Medicine, Radiotherapy & Molecular Imaging (AREA)
- Optics & Photonics (AREA)
- Chemical & Material Sciences (AREA)
- Crystallography & Structural Chemistry (AREA)
- Measurement Of Radiation (AREA)
Abstract
The invention discloses an X-ray detection array pixel unit, a manufacturing process of a double-layer scintillator array and a double-layer energy spectrum CT detector based on the pixel unit. The double-layer scintillator array adopts a two-dimensional array processing technology, and the integration technology of the detector submodule adopts the same superposition assembly technology as that of a conventional CT detector, so that the sensitive area of the detector is not influenced, the imaging dosage efficiency can be effectively maintained, and the cost of the detector can be controlled.
Description
Technical Field
The invention relates to the field of semiconductor photoelectric detectors, in particular to an X-ray detection array pixel unit, a double-layer scintillator array manufacturing process and a double-layer energy spectrum CT detector.
Background
The X-ray response signal acquired in conventional CT is proportional to the energy integral of all the different energy X-rays detected by the detector, without any spectral information of the X-rays. In recent years, the rapidly developed energy spectrum CT is obtained, more image information than the conventional CT is obtained by collecting and distinguishing response signals of X-rays with different energies, the possibility is provided for carrying out density resolution optimization on images according to energy weights and realizing substance resolution of imaging objects, and the image quality, the dose efficiency and the accuracy of clinical diagnosis are obviously improved on the basis of the conventional CT. The technology and application research of energy spectrum CT has become an important development direction of CT medical imaging technology.
A more sophisticated method of acquiring energy spectrum data in CT applications is to use a dual layer scintillator X-ray detector instead of the single layer scintillator X-ray detector in conventional CT. The detector array in conventional CT is typically formed by stitching multiple detector modules along the X-direction, each of which is precisely positioned by a detector rail machined into an arc, ensuring that the distance between each module and the focal point of the bulb is the same, as shown in fig. 1. The detector module is generally formed by arranging a plurality of detector minimum sub-modules along a Z direction perpendicular to an XY plane, and one structure of the detector minimum sub-modules is shown in fig. 2 and is composed of a collimator array, a scintillator array, a photosensitive array, an analog-to-digital conversion chip, a mounting block for accurate assembly and positioning and the like. The scintillator array and the photosensitive array are main devices for signal conversion and consist of two-dimensional pixel arrays. The structure of a conventional CT detector signal conversion pixel unit is shown in FIG. 3, wherein a scintillator pixel converts incident X-rays into visible light, a photosensitive array pixel converts visible light generated by the scintillator pixel into an electrical signal, and the electrical signal is transmitted to a subsequent analog-to-digital conversion circuit through signal connection of a substrate. The traditional CT detector adopts a single-layer scintillator array and an integral signal acquisition mode, and cannot distinguish energy spectrum information of incident X rays. Compared with the structure of the traditional CT detector, the existing double-layer energy spectrum CT detector mainly has the difference of a more complex signal conversion pixel structure, as shown in fig. 4, each signal conversion pixel consists of two laminated scintillator sub-pixels and two side-coupled photosensitive array sub-pixels, and the top layer and the bottom layer of the scintillator sub-pixels are connected by a light reflection layer and realize the isolation of light signals. The top-layer scintillator sub-pixels mainly detect low-energy parts in incident X-rays, the bottom-layer scintillator sub-pixels mainly detect high-energy parts in the incident X-rays, and energy spectrum information corresponding to the high-energy and low-energy parts of the incident X-rays can be obtained in an imaging process by simultaneously reading signals of the two sub-pixels.
The existing double-layer energy spectrum CT detector adopts a side coupling design between a scintillator sub-pixel and a photosensitive array sub-pixel, so that two main defects of the detector which cannot be avoided are determined: (1) The detector unit is formed by three-dimensional precise integration and assembly of physically divided scintillator pixels and photosensitive array pixels, and the complex processing and assembly process is not beneficial to mass production and cost control of the detector; (2) In the incident direction of the X-ray, the photosensitive array pixel occupies a certain sensitive area of the detector, thereby affecting the imaging dosage efficiency.
Disclosure of Invention
The technical purpose is that: aiming at the defects in the prior art, the invention discloses an X-ray detection array pixel unit and a double-layer energy spectrum CT detector, which have novel signal conversion pixel structure design, use a two-dimensional array processing technology of a double-layer scintillator array and use a two-dimensional area scintillator array and photosensitive array superposition assembly technology of a conventional CT detector to realize energy spectrum signal reading of the double-layer detector, thereby effectively avoiding the main defects of the existing double-layer energy spectrum CT detector.
The technical scheme is as follows: in order to achieve the technical purpose, the invention adopts the following technical scheme:
An X-ray detection array pixel unit, characterized in that: the light source comprises a top layer scintillator pixel, a bottom layer scintillator pixel, a thin film light filter layer and a photosensitive array pixel which are sequentially arranged in the light incidence direction;
The top-layer scintillator pixels and the bottom-layer scintillator pixels form a double-layer scintillator array, and are made of scintillator materials with obviously different luminous spectrums and clearly distinguishable, and are used for converting incident X rays into visible light; the light signals emitted by the top-layer scintillator pixels enter the bottom-layer scintillator pixels through the contact surface, and the bottom surfaces of the bottom-layer scintillator pixels are light output surfaces; the outer parts of the top scintillator pixels and the bottom scintillator pixels are provided with light reflecting layers;
The film light filter layer comprises two filter areas which are arranged on the same layer, namely a first filter area of the film light filter layer and a second filter area of the film light filter layer, and the filter areas correspond to the luminous spectrums of the top-layer scintillator and the bottom-layer scintillator respectively;
the photosensitive array pixels are used for converting visible light generated by the scintillator pixels into electric signals, each photosensitive array pixel comprises a first photosensitive sub-pixel corresponding to a first filtering area of the thin film light filtering layer and a second photosensitive sub-pixel corresponding to a second filtering area of the thin film light filtering layer, and output signals of the first photosensitive sub-pixel and the second photosensitive sub-pixel are respectively corresponding to X-ray responses of the top-layer scintillator and the bottom-layer scintillator.
Preferably, the first filter area of the thin film light filtration layer is equal to the first photosensitive array sub-pixel in size, and the second filter area is equal to the second photosensitive array sub-pixel in size.
Preferably, the top layer scintillator pixels and the bottom layer scintillator pixels are connected in a direct optical coupling manner, so that light emission of the top layer scintillator pixels can enter the bottom layer scintillator pixels.
Preferably, the combination of the top layer scintillator pixels and the bottom layer scintillator pixels adopts any one of ZnSe/GOS, baF2/CdWO4 and ZnSe/YAG.
Preferably, the combination of the top-layer scintillator pixels and the bottom-layer scintillator pixels adopts ZnSe/GOS, and the thicknesses of the top-layer scintillator pixels and the top-layer scintillator pixels are respectively 0.5mm and 1.0mm.
Preferably, the first filtering area of the thin film light filtering layer is a short-pass filtering layer and only allows photons with the wavelength greater than or equal to 550nm to pass through, and the second filtering area of the thin film light filtering layer is a long-pass filtering layer and only allows photons with the wavelength greater than 550nm to pass through.
Preferably, the area ratio of the first filtering area of the thin film light filtering layer to the second filtering area of the thin film light filtering layer is 1:1.
A process for manufacturing a two-layer scintillator array, characterized by: the following steps are sequentially performed:
(1) Adopting a top-layer scintillator wafer and a bottom-layer scintillator wafer as raw materials, respectively polishing, and bonding the top-layer scintillator and the bottom-layer scintillator wafer through optical coupling glue to obtain an integrated double-layer scintillator wafer;
(2) Cutting a plurality of two-dimensional separation grooves from one end of a top-layer scintillator wafer of the double-layer scintillator wafer, wherein the depth of each two-dimensional separation groove is larger than the design thickness of the double-layer scintillator array and smaller than the actual thickness of the double-layer scintillator wafer, and the spacing between adjacent pixels separated by each two-dimensional separation groove is equal to the required size of the pixels of the double-layer scintillator array;
(3) Filling a light reflection layer material from one end of a top scintillator wafer of the double-layer scintillator wafer and curing to form a light reflection layer covering the top scintillator wafer and separating adjacent scintillator pixels;
(4) Grinding the double-layer scintillator wafer with the light reflection layer process from one end of the bottom-layer scintillator wafer until the thickness of the bottom-layer scintillator wafer is reduced to a preset thickness, and polishing from one end of the bottom-layer scintillator wafer to form a coupling surface with the thin-film light filter layer; cutting along the two-dimensional separation groove to obtain the required double-layer scintillator.
Preferably, the areas of the top scintillator wafer and the bottom scintillator wafer in the step (1) are determined by wafer raw materials and production processes;
The design thickness of the top scintillator wafer in the step (2) is the design thickness of the top scintillator pixel, and the thickness of the bottom scintillator wafer is larger than the design thickness of the bottom scintillator pixel;
in the step (4), the top end of the top scintillator wafer and the two-dimensional separation groove are covered and filled by the light reflection layer, and the bottom surface of the bottom scintillator wafer is the light output surface of the pixel.
A dual-layer energy spectrum CT detector, characterized by: the X-ray detector comprises a plurality of double-layer energy spectrum CT detector minimum sub-modules, wherein each minimum sub-module comprises a plurality of X-ray detection array pixel units, and a top-layer scintillator array pixel is used as an incident end of X-rays.
The beneficial effects are that: compared with the existing double-layer CT energy spectrum detector technology, the signal conversion pixel structure and the double-layer CT energy spectrum detector have the following technical characteristics and effects:
(1) The double-layer scintillator array is processed by two different scintillator materials, and under the irradiation of X rays, the two scintillator materials have different luminous spectrums which can be obviously distinguished and can be distinguished by a color filtering method; the light emitted by the top layer scintillator sub-pixels can penetrate the bottom layer scintillator sub-pixels to reach the light output surface by adopting a direct optical coupling mode between the top layer and the bottom layer scintillator sub-pixels.
(2) A thin film light filter layer is arranged between the scintillator pixels and the photosensitive array pixels and corresponds to the light output surface of each scintillator pixel, the thin film light filter layer is divided into two different areas and corresponds to the two photosensitive array sub-pixels respectively; the light filtering design of one area only allows the light emitting spectrum of the top scintillator to pass through, and the light filtering design of the other area only allows the light emitting spectrum of the bottom scintillator to pass through, so that the output signals of the two photosensitive array sub-pixels respectively correspond to the X-ray responses of the top scintillator sub-pixels and the bottom scintillator sub-pixels, and the energy spectrum information of the incident X-rays is effectively provided.
(3) The existing double-layer CT detector adopts a scintillator array and a photosensitive array which are physically divided, and realizes signal readout of the double-layer detector through a complex three-dimensional integration process, and the processing cost is obviously higher than that of a two-dimensional array plane superposition assembly process adopted by a conventional CT detector; the double-layer CT detector adopts a two-dimensional continuous scintillator array, a photosensitive array and a pixelated film light filter layer, the double-layer scintillator array adopts a two-dimensional array processing technology, and the integration technology adopts the same plane superposition assembly technology as that of the conventional CT detector, so that the processing cost similar to that of the conventional CT detector can be effectively maintained.
(4) The existing double-layer CT detector adopts a scintillator array and photosensitive array side coupling mode, so that the photosensitive array occupies a certain detector sensitivity area in the X-ray incidence direction, thereby affecting the imaging dosage efficiency; the double-layer CT detector disclosed by the invention adopts a two-dimensional area scintillator array and photosensitive array superposition assembly process used by a conventional CT detector, the sensitive area of the detector is not influenced, and the imaging dosage efficiency can be effectively maintained.
Drawings
FIG. 1 is a schematic view of a CT detector array;
FIG. 2 is a schematic diagram of an embodiment of a detector minimum sub-module;
FIG. 3 is a schematic diagram of a conventional single-layer CT detector signal-converting pixel;
FIG. 4 is a schematic diagram of a signal conversion pixel of a conventional dual-layer CT spectrum detector;
FIG. 5 is a schematic diagram of a signal conversion pixel of a dual-layer CT spectrum detector according to the present invention;
FIG. 6 is a luminescence spectrum of GOS and ZnSe scintillators;
FIG. 7 is the transparency of GOS scintillators at different wavelengths;
FIG. 8 is a schematic diagram of a fabrication process for a dual layer scintillator array;
fig. 9 is a schematic cross-sectional view of a 16x16 pixel dual layer scintillator array.
Detailed Description
The invention provides a double-layer energy spectrum CT detector, which is formed by processing a two-dimensional continuous double-layer scintillator array, a photosensitive array and a pixelated film light filter layer in the pixel structure design of the double-layer CT detector.
The dual layer scintillator array is fabricated using different top and bottom layer scintillator materials according to the process steps shown in fig. 8. Under the irradiation of X-rays, the scintillator materials of the top-layer and bottom-layer scintillator sub-pixels have different luminous spectrums which can be distinguished obviously, and can be distinguished by a color filtering method. The light emitted by the top layer scintillator sub-pixels can penetrate the bottom layer scintillator sub-pixels to reach the light output surface by adopting a direct optical coupling mode between the top layer and the bottom layer scintillator sub-pixels. A thin film light filter layer is arranged between the scintillator pixels and the photosensitive array pixels, corresponds to the light output surface of each scintillator pixel, and is divided into two different areas, wherein the light filter design of one area only allows the light-emitting spectrum of the top scintillator to pass through, and the light filter design of the other area only allows the light-emitting spectrum of the bottom scintillator to pass through.
As an example of the thin film light filtration layer, the function thereof may be achieved by preparing a plurality of light transmission films having different refractive indexes on a flexible or rigid transparent substrate, and by adjusting the thickness of each light transmission film, the light filtration effect of a specific wavelength range may be achieved. Light-transmitting thin film materials of different refractive indices include, but are not limited to, various combinations of Ag and SiO2, ag and TiO2, and the like.
The photosensitive array pixel is divided into two independently read-out sub-pixels, the areas of the two sub-pixels respectively correspond to the two areas of the thin film light filtering layer, and output signals of the two photosensitive array sub-pixels are transmitted to two channels of a subsequent analog-digital conversion circuit through substrate connection and are read out simultaneously.
In the present invention, the dual-layer scintillator array can be manufactured by using the processing steps shown in fig. 8, including the steps of:
(1) After polishing the top scintillator and the bottom scintillator wafers, firmly and directly coupling light through light coupling glue to obtain an integrated wafer of the double-layer scintillator; the area of the wafer is determined by the wafer raw material and the production process and can be much larger than the size of the double-layer scintillator array used in the minimum submodule of the final detector; the thickness of the top scintillator wafer is the design thickness of the top scintillator pixels, and the thickness of the bottom scintillator wafer is larger than the design thickness of the bottom scintillator pixels, so that the required machining allowance is provided for the subsequent machining steps;
(2) Cutting and processing two-dimensional separation grooves between adjacent pixels from one end of a top-layer scintillator on a double-layer scintillator wafer according to the design thickness of a light reflection layer between the pixels in the scintillator array, wherein the depth of the grooves is controlled to be larger than the design thickness of the double-layer scintillator array and smaller than the thickness of the double-layer scintillator wafer, so that the whole wafer is still connected by the bottom-layer scintillator wafer which is not cut, and an integrated structure is maintained;
(3) Filling a light reflection layer material from one end of the top-layer scintillator according to design and curing to form a light reflection layer covering the top-layer scintillator and between adjacent pixels;
(4) And grinding the wafer with the light reflection layer process from one end of the bottom layer scintillator until the thickness of the bottom layer scintillator is reduced to the design thickness of the scintillator pixels, and polishing from one end of the bottom layer scintillator to form a coupling surface with the light filtering layer. So far, the structural processing of the double-layer scintillator array has been completed, which can be cut into the required array size in the detector minimum sub-module for use.
As an example, fig. 9 is a schematic cross-sectional view of a cut 16x16 pixel dual-layer scintillator array, each dual-layer scintillator pixel being composed of a top-layer scintillator pixel directly optically coupled to a bottom-layer scintillator pixel, the top ends of the top-layer scintillator pixels and between adjacent scintillator pixels being covered and separated by a light reflective layer, the bottom surface of the bottom-layer scintillator pixel being the light output surface of the pixel.
As shown in fig. 5, which is a specific embodiment of the pixel structure of the present invention, znSe (zinc selenide) and GOS (gadolinium oxysulfide) scintillator materials are used as the top and bottom scintillator sub-pixels, respectively, with thicknesses of 0.5mm and 1.0mm, respectively. The luminescence spectra of the two scintillators under X-ray irradiation are shown in fig. 6, have different main peak wavelengths, and are clearly discriminated in spectrum. The transparency curve of a GOS scintillator 1.0mm thick is shown in fig. 7, and the transparency is good in the emission spectrum region of the ZnSe scintillator, so that most of the emission photons of the ZnSe scintillator can pass through the GOS scintillator to reach the coupling surface with the light filter layer.
The two areas of the light filter layer are respectively designed into a short-pass filter layer with the cut-off wavelength of 550nm, and only photons with the wavelength less than 550nm are allowed to pass through; and a long pass filter layer allowing only photons having a wavelength greater than 550nm to pass. Thus, the photosensitive array sub-pixel corresponding to the short-pass filter region reads out signals from the GOS scintillator sub-pixel, and the photosensitive array sub-pixel corresponding to the long-pass filter region reads out signals from the ZnSe scintillator sub-pixel, and information about the incident X-ray energy spectrum can be obtained by simultaneously reading out the two signals.
The practical application of the present invention is not limited to the combination of the pair of scintillator materials ZnSe/GOS, but any combination of scintillator materials satisfying the above-described characteristics is within the scope of the present invention, such as the combination of scintillator materials BaF2/GOS, baF2/CdWO4, znSe/YAG, etc. In addition, in the pixel structure shown in fig. 5, the two areas of the light filtering layer and the corresponding two photosensitive array sub-pixel areas are divided according to a proportion of 50% -50%, and in practical application, the proportion can be optimized and adjusted according to the relative size of signals.
The foregoing is only a preferred embodiment of the invention, it being noted that: it will be apparent to those skilled in the art that various modifications and adaptations can be made without departing from the principles of the present invention, and such modifications and adaptations are intended to be comprehended within the scope of the invention.
Claims (9)
1. An X-ray detection array pixel unit, characterized in that: a top layer scintillator pixel, a bottom layer scintillator pixel, a thin film light filter layer and a photosensitive array pixel are sequentially arranged in the light incidence direction;
the top-layer scintillator pixels and the bottom-layer scintillator pixels form a double-layer scintillator array, and are made of scintillator materials with different luminous spectrums and used for converting incident X-rays into visible light; the outer parts of the top scintillator pixels and the bottom scintillator pixels are provided with light reflecting layers;
the thin film light filter layer comprises two filter areas which are arranged on the same layer, namely a first filter area of the thin film light filter layer and a second filter area of the thin film light filter layer, and the two filter areas correspond to the luminous spectrum areas of the top scintillator pixels and the bottom scintillator pixels respectively;
The photosensitive array pixels are used for converting visible light generated by the scintillator pixels into electric signals, each photosensitive array pixel comprises a first photosensitive sub-pixel corresponding to a first filtering area of the thin film light filtering layer and a second photosensitive sub-pixel corresponding to a second filtering area of the thin film light filtering layer, and output signals of the first photosensitive sub-pixel and the second photosensitive sub-pixel respectively correspond to X-ray responses of the top-layer scintillator and the bottom-layer scintillator;
the processing technology of the double-layer scintillator array comprises the following steps:
(1) After polishing the top scintillator and the bottom scintillator wafers, firmly and directly coupling light through light coupling glue to obtain an integrated wafer of the double-layer scintillator; the area of the wafer is determined by the wafer raw material and the production process and can be much larger than the size of the double-layer scintillator array used in the minimum submodule of the final detector; the thickness of the top scintillator wafer is the design thickness of the top scintillator pixels, and the thickness of the bottom scintillator wafer is larger than the design thickness of the bottom scintillator pixels, so that the required machining allowance is provided for the subsequent machining steps;
(2) Cutting and processing two-dimensional separation grooves between adjacent pixels from one end of a top-layer scintillator on a double-layer scintillator wafer according to the design thickness of a light reflection layer between the pixels in the scintillator array, wherein the depth of the grooves is controlled to be larger than the design thickness of the double-layer scintillator array and smaller than the thickness of the double-layer scintillator wafer, so that the whole wafer is still connected by the bottom-layer scintillator wafer which is not cut, and an integrated structure is maintained;
(3) Filling a light reflection layer material from one end of the top-layer scintillator according to design and curing to form a light reflection layer covering the top-layer scintillator and between adjacent pixels;
(4) Grinding a wafer with the light reflection layer process from one end of a bottom layer scintillator until the thickness of the bottom layer scintillator is reduced to the design thickness of a scintillator pixel, and polishing from one end of the bottom layer scintillator to form a coupling surface with a light filtering layer; so far, the structural processing of the double-layer scintillator array has been completed, which can be cut into the required array size in the detector minimum sub-module for use.
2. The X-ray detection array pixel unit of claim 1, wherein: the first filter area of the film light filter layer is equal to the first photosensitive array sub-pixel in size, and the second filter area of the film light filter layer is equal to the second photosensitive array sub-pixel in size.
3. The X-ray detection array pixel unit of claim 1, wherein: the top layer scintillator pixels and the bottom layer scintillator pixels are connected in a direct optical coupling mode.
4. The X-ray detection array pixel unit of claim 1, wherein: the combination of the top-layer scintillator pixels and the bottom-layer scintillator pixels adopts any one of ZnSe/GOS, baF2/CdWO4 and ZnSe/YAG.
5. The X-ray detection array pixel unit of claim 1, wherein: the combination of the top-layer scintillator pixels and the bottom-layer scintillator pixels adopts ZnSe/GOS, and the thicknesses of the top-layer scintillator pixels and the top-layer scintillator pixels are respectively 0.5mm and 1.0mm.
6. The X-ray detection array pixel unit of claim 5, wherein: the first filtering area of the thin film light filtering layer is a short-pass filtering layer and only allows photons with the wavelength greater than or equal to 550nm to pass through, and the second filtering area of the thin film light filtering layer is a long-pass filtering layer and only allows photons with the wavelength greater than 550nm to pass through.
7. The X-ray detection array pixel unit of claim 1, wherein: the area ratio of the first filtering area of the film light filtering layer to the second filtering area of the film light filtering layer is 1:1.
8. The X-ray detection array pixel unit of claim 1, wherein: the areas of the top scintillator wafer and the bottom scintillator wafer in the step (1) are determined by wafer raw materials and production processes;
The design thickness of the top scintillator wafer in the step (2) is the design thickness of the top scintillator pixel, and the thickness of the bottom scintillator wafer is larger than the design thickness of the bottom scintillator pixel;
in the step (4), the top end of the top scintillator wafer and the two-dimensional separation groove are covered and filled by the light reflection layer, and the bottom surface of the bottom scintillator wafer is the light output surface of the pixel.
9. A dual-layer energy spectrum CT detector, characterized by: a plurality of dual-layer energy spectrum CT detector minimum sub-modules, each comprising a plurality of X-ray detection array pixel units as defined in any one of claims 1-8, with top-layer scintillator array pixels as the X-ray incident end.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CN202111030092.4A CN113671553B (en) | 2021-09-03 | 2021-09-03 | X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector |
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CN202111030092.4A CN113671553B (en) | 2021-09-03 | 2021-09-03 | X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector |
Publications (2)
| Publication Number | Publication Date |
|---|---|
| CN113671553A CN113671553A (en) | 2021-11-19 |
| CN113671553B true CN113671553B (en) | 2024-05-31 |
Family
ID=78548357
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CN202111030092.4A Active CN113671553B (en) | 2021-09-03 | 2021-09-03 | X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector |
Country Status (1)
| Country | Link |
|---|---|
| CN (1) | CN113671553B (en) |
Families Citing this family (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CN115052116B (en) * | 2022-06-14 | 2023-09-19 | 无锡鉴微华芯科技有限公司 | Thick silicon pixel detector with non-uniformly distributed pixel units and preparation method thereof |
Citations (7)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CN101937095A (en) * | 2009-06-30 | 2011-01-05 | 同方威视技术股份有限公司 | Dual energy X ray detector and dual energy X ray detector array device |
| CN104285162A (en) * | 2012-05-07 | 2015-01-14 | 皇家飞利浦有限公司 | Multi-layer horizontal computed tomography (CT) detector array with at least one thin photosensor array layer disposed between at least two scintillator array layers |
| CN105723243A (en) * | 2013-11-15 | 2016-06-29 | 皇家飞利浦有限公司 | Double-sided organic photodetector on flexible substrate |
| CN110168406A (en) * | 2017-01-02 | 2019-08-23 | 皇家飞利浦有限公司 | X-ray detector and x-ray imaging device |
| CN211066679U (en) * | 2019-10-31 | 2020-07-24 | 南京安科医疗科技有限公司 | CT detector heat abstractor |
| CN112022179A (en) * | 2020-09-03 | 2020-12-04 | 南京安科医疗科技有限公司 | Detector unit arrangement method of high-row-number CT detector |
| CN112558134A (en) * | 2021-03-01 | 2021-03-26 | 同源微（北京）半导体技术有限公司 | Hybrid X-ray detector |
Family Cites Families (5)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US7054408B2 (en) * | 2003-04-30 | 2006-05-30 | General Electric Company | CT detector array having non pixelated scintillator array |
| RU2487373C2 (en) * | 2007-12-21 | 2013-07-10 | Конинклейке Филипс Электроникс Н.В. | Radiation-sensitive detector with scintillator in composite resin |
| WO2010058309A2 (en) * | 2008-11-18 | 2010-05-27 | Koninklijke Philips Electronics N.V. | Spectral imaging detector |
| WO2011148276A2 (en) * | 2010-05-24 | 2011-12-01 | Koninklijke Philips Electronics N.V. | Ct detector including multi-layer fluorescent tape scintillator with switchable spectral sensitivity |
| US9335425B2 (en) * | 2011-10-12 | 2016-05-10 | Koninklijke Philips N.V. | Modelling of ToF-DOI detector arrays |
-
2021
- 2021-09-03 CN CN202111030092.4A patent/CN113671553B/en active Active
Patent Citations (7)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CN101937095A (en) * | 2009-06-30 | 2011-01-05 | 同方威视技术股份有限公司 | Dual energy X ray detector and dual energy X ray detector array device |
| CN104285162A (en) * | 2012-05-07 | 2015-01-14 | 皇家飞利浦有限公司 | Multi-layer horizontal computed tomography (CT) detector array with at least one thin photosensor array layer disposed between at least two scintillator array layers |
| CN105723243A (en) * | 2013-11-15 | 2016-06-29 | 皇家飞利浦有限公司 | Double-sided organic photodetector on flexible substrate |
| CN110168406A (en) * | 2017-01-02 | 2019-08-23 | 皇家飞利浦有限公司 | X-ray detector and x-ray imaging device |
| CN211066679U (en) * | 2019-10-31 | 2020-07-24 | 南京安科医疗科技有限公司 | CT detector heat abstractor |
| CN112022179A (en) * | 2020-09-03 | 2020-12-04 | 南京安科医疗科技有限公司 | Detector unit arrangement method of high-row-number CT detector |
| CN112558134A (en) * | 2021-03-01 | 2021-03-26 | 同源微（北京）半导体技术有限公司 | Hybrid X-ray detector |
Non-Patent Citations (1)
| Title |
|---|
| 轮胎U型探测器开发与应用研究;宋云峰;;科技风;20120325(06);21-22 * |
Also Published As
| Publication number | Publication date |
|---|---|
| CN113671553A (en) | 2021-11-19 |
Similar Documents
| Publication | Publication Date | Title |
|---|---|---|
| RU2386981C2 (en) | Improved detector matrix for spectral computed tomography | |
| EP1876955B1 (en) | Double decker detector for spectral ct | |
| RU2603240C2 (en) | Radiation detector | |
| US4982096A (en) | Multi-element radiation detector | |
| KR101022150B1 (en) | Radiation image converting panel, scintillator panel and radiation image sensor | |
| RU2542588C2 (en) | Measurement system of image generation with print matrix of photodetectors | |
| US9000382B2 (en) | Spectral imaging detector | |
| US8481952B2 (en) | Scintillation separator | |
| RU2488141C2 (en) | Scintillator for imaging device, scintillator module, imaging device with said scintillator and method of manufacturing scintillator | |
| US7399972B2 (en) | Component for radiation detector and radiation detector | |
| JPS6211313B2 (en) | ||
| JPS61110079A (en) | Radiation detector | |
| JP2008510131A (en) | Arrangement of scintillator and anti-scatter grid | |
| CN113671553B (en) | X-ray detection array pixel unit, manufacturing process and double-layer energy spectrum CT detector | |
| US6823038B2 (en) | X-ray detector array and method for manufacturing same | |
| US6793857B2 (en) | Method for producing a one- or multidimensional detector array | |
| JP2720159B2 (en) | Multi-element radiation detector and manufacturing method thereof | |
| WO2007113899A1 (en) | Radiation detector | |
| JP2594966B2 (en) | Radiation detecting element block and method of manufacturing the same | |
| JP2003084067A (en) | X-ray detector, and x-ray ct device using the same | |
| JPH1184013A (en) | Radiation detector | |
| JPH1010235A (en) | Device for detecting multi-element radiation and its manufacture | |
| KR101405536B1 (en) | Directly-Deposited Scintillator Panel and Manufacturing Method thereof | |
| JP2019163970A (en) | Scintillator array, radiation detector, and radiation computer tomography apparatus | |
| CN115902991A (en) | PET detector based on partially continuous crystals and PET imaging system |
Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| PB01 | Publication | ||
| PB01 | Publication | ||
| SE01 | Entry into force of request for substantive examination | ||
| SE01 | Entry into force of request for substantive examination | ||
| GR01 | Patent grant | ||
| GR01 | Patent grant |