Photon counting detector

Abstract

The embodiment of the invention provides a photon counting detector, and the detector comprises a scintillation crystal array which is used for converting X-rays into fluorescence; the photoelectric converter is used for converting the fluorescence into an electronic signal; the photon counting circuit is used for processing the electronic signal to obtain photon counting data; the control reading circuit is used for reading the photon counting data and transmitting the photon counting data to the external reading circuit; the scintillation crystal array is made of scintillation crystal materials, and the luminescence decay time of the scintillation crystal materials is within 25 ns.

Description

TECHNICAL FIELD

[0001] The present specification relates to the field of photon detection, and in particular, to a photon counting detector. BACKGROUND

[0002] Direct conversion detectors made of room temperature semiconductor materials (e.g., cadmium telluride, cadmium zinc telluride, monocrystalline silicon, etc.) are the mainstream detector type of today's photon counting CT (PCCT, Photon Counting Computed Tomography) devices. The direct conversion detector can realize single photon measurement, but the manufacturing process of such detector material is complex, the yield is low, resulting in high cost, and the polarization effect under high flux causes the detector to work unstably, and there are problems such as poor single photon energy resolution and missing photon counting.

[0003] Therefore, it is desirable to provide a photon counting detector to improve the above shortcomings of the direct conversion detector to realize accurate single photon measurement. SUMMARY

[0004] One or more embodiments of the present specification provide a photon counting detector. The photon counting detector includes: a scintillation crystal array for converting X-rays into fluorescent light; a photoelectric converter for converting the fluorescent light into an electronic signal; a photon counting circuit for processing the electronic signal to obtain photon counting data; and a control readout circuit for reading out the photon counting data and transmitting the photon counting data to an external readout circuit. The scintillation crystal array is made of a scintillation crystal material, and the light emission decay time of the scintillation crystal material is within 25 ns.

[0005] In some embodiments, the scintillation crystal array includes a plurality of scintillation crystal units, and the aspect ratio of the scintillation crystal array is greater than 2, where the aspect ratio is the ratio of the thickness of the scintillation crystal unit to the spacing of adjacent scintillation crystal units.

[0006] In some embodiments, the photoelectric converter includes a pixel array, and the pixel array includes a plurality of pixels, wherein each pixel includes a photodiode, and the operating voltage of the photodiode can be configured; and the fill factor of each of the plurality of pixels is greater than 50%, where the fill factor is the ratio of the area of the light-sensitive surface of the photodiode to the area of the pixel.

[0007] In some embodiments, the processing includes amplitude discrimination and photon counting, the photon counting data includes a plurality of photon counting sub-data, the photon counting circuit includes a plurality of signal processing circuits, a plurality of threshold comparator groups and a plurality of counter groups, wherein the plurality of signal processing circuits are respectively connected to a plurality of pixels in the pixel array, and each signal processing circuit includes a preamplifier and a filter; the plurality of threshold comparator groups are respectively connected to the plurality of signal processing circuits, and each of the plurality of threshold comparator groups includes a plurality of threshold comparators, wherein each threshold comparator is configured to perform the amplitude discrimination, including: determining whether the amplitude of the electronic signal is higher than a preset amplitude threshold of the threshold comparator; and generating a digital pulse in response to the amplitude of the electronic signal being higher than the amplitude threshold of the threshold comparator; the plurality of counter groups are respectively connected to the plurality of threshold comparator groups, and each of the plurality of counter groups includes a plurality of counters, wherein the plurality of counters in each counter group are respectively connected to the plurality of threshold comparators in a threshold comparator group connected to the counter group, and each counter is configured to perform photon counting, including: counting the digital pulses generated by the connected threshold comparators to determine the photon counting sub-data.

[0008] In some embodiments, the pixel array includes a plurality of pixel blocks, and each pixel block includes a plurality of pixels; the readout control circuit reads out the photon counting sub-data corresponding to each pixel, or reads out the photon counting sub-zone data corresponding to each pixel block.

[0009] In some embodiments, the photon counting detector further includes a fluorescent transmission structure, and the fluorescent transmission structure includes a reflective layer, a light coupling layer and a microlens layer, wherein the reflective layer is located on the front surface and the side surface of each scintillation crystal unit, the front surface is a surface of the scintillation crystal unit facing the X-ray source, and the side surface is parallel to the incident direction of the X-ray; the light coupling layer and the microlens layer are located between the scintillation crystal array and the photoelectric converter.

[0010] In some embodiments, the photon counting detector further includes an anti-scattering grating, and the anti-scattering grating is located on the side of the scintillation crystal array facing the X-ray source, the anti-scattering grating includes a grating wall and a plurality of transmission holes, wherein the pitch of adjacent transmission holes is equal to the pitch of adjacent pixels in the pixel array; the thickness of the grating wall is less than 150 μm.

[0011] In some embodiments, the scintillation crystal array is directly generated on the surface of the photoelectric converter, and the photoelectric converter and the photon counting circuit are made by an integrated process.

[0012] In some embodiments, an output interface of the photon counting circuit is in communication with a packaging electrode in the control readout circuit to realize signal transmission between the photon counting circuit and the control readout circuit, and the output interface is made by a through silicon via process.

[0013] The one or more embodiments of the specification provide a scanning imaging device. The scanning imaging device comprises an X-ray source and a photon counting detector provided by the one or more embodiments of the specification. BRIEF DESCRIPTION OF DRAWINGS

[0014] The specification will be further illustrated in the way of exemplary embodiments, which will be described in detail with reference to the accompanying drawings. These embodiments are not restrictive, and in these embodiments, the same numbers represent the same structures, wherein:

[0015] Figure 1 is a schematic diagram of an application scenario of a scanning imaging device according to some embodiments of the specification; Figure 2 is a schematic diagram of imaging principle of a scanning imaging device according to some embodiments of the specification; Figure 3 is an exemplary structural diagram of a photon counting detector according to some embodiments of the specification; Figure 4 is a schematic diagram of fluorescence conversion principle of a photon counting detector according to some embodiments of the specification; Figure 5 is a schematic diagram of circuit structure of a photon counting circuit according to some embodiments of the specification; Figure 6 is a schematic diagram of circuit structure of a photon counting circuit according to some other embodiments of the specification; and Figure 7 is a schematic diagram of structure of an integrated photoelectric converter and photon counting circuit according to some embodiments of the specification. DETAILED DESCRIPTION

[0016] In order to more clearly illustrate the technical solutions of the embodiments of the specification, the following will briefly introduce the drawings needed to be used in the embodiment description. Obviously, the drawings in the following description are only some examples or embodiments of the specification, and for those skilled in the art, the specification can be applied to other similar scenarios without creative labor. Unless it is obvious from the language environment or otherwise stated, the same reference numbers in the drawings represent the same structure or operation.

[0017] It should be understood that the "system", "device", "unit" and / or "module" used herein is a method for distinguishing different components, elements, parts, portions or assemblies at different levels. However, if other words can achieve the same purpose, the words can be replaced by other expressions.

[0018] As shown in the specification and claims herein, unless the context clearly indicates otherwise, the words "comprise", "comprising", "consist of", "consisting of", "include", "including", "involving", "contain", "containing" or "characterized by" are not used to convey an exclusive inclusion of the steps or elements specified, but rather to convey an inclusion of those specified steps or elements as well as additional steps or elements not specified.

[0019] Flow diagrams are used in the description of the specification to demonstrate the operation of the systems according to embodiments of the specification. It should be understood that the operations in the front or back are not necessarily performed in order. Instead, various steps can be processed in different order or simultaneously. At the same time, other operations can be added to these processes, or one or more steps can be removed from these processes.

[0020] The detector of the mainstream photon counting CT (PCCT) device today includes a direct conversion photon counting detector made of room temperature semiconductor material. The direct conversion photon counting detector can realize single photon measurement, complete energy identification of X-rays, and thus can provide excellent spatial resolution without affecting the count rate and spectral resolution, but the detector material manufacturing process is complex, resulting in high cost, and at high flux, the polarization effect of the detector leads to unstable signal output of the detector, reduces the energy resolution and increases the image noise, in addition, the excited charge clusters in the detector will be collected by the electrodes of adjacent pixels due to diffusion effect, resulting in problems such as single photon energy resolution degradation and photon counting omission.

[0021] To improve the above shortcomings of the existing detector, some embodiments of the specification propose a photon counting detector, which can improve the single photon energy resolution on the basis of realizing single photon measurement, avoid photon counting omission, and thus reduce image noise and reduce the cost of the detector.

[0022] Figure 1 is a schematic diagram of the application scenario of the scanning imaging device according to some embodiments of the specification. In some embodiments, as shown in Figure 1 the application scenario 100 of the scanning imaging device (hereinafter referred to as the application scenario 100) can include a processing device 110, a network 120, a storage device 130, a scanning imaging device 140 and a terminal device 150.

[0023] The processing device 110 refers to a device for processing data and / or information from at least one component of the application scenario 100 or an external data source (e.g., a cloud data center). The processing device 110 can access data or information from the storage device 130, the scanning imaging device 140, and / or the terminal device 150 through the network 120. The processing device 110 can also directly connect to the storage device 130, the scanning imaging device 140, and / or the terminal device 150 to access information and / or data. For example, the processing device 110 can obtain photon counting data output by a photon counting detector (not shown in the figure) in the scanning imaging device 140 from the scanning imaging device 140. For another example, the processing device 110 can determine a concentration map of a base substance based on the above-mentioned photon counting data, thereby further generating a scanning image of a scanning object (e.g., a patient scanned by the scanning imaging device 140).

[0024] The network 120 can connect various components of the application scenario 100 and / or connect the application scenario 100 with external resource parts. In some embodiments, one or more components (e.g., the processing device 110, the storage device 130, the scanning imaging device 140, and / or the terminal device 150) of the application scenario 100 can exchange information and / or data through the network 120.

[0025] In some embodiments, the network 120 can be a wired network or a wireless network, or a combination of the two. In some embodiments, the network 120 can include one or more network access points. For example, the network 120 can include wired or wireless network access points (e.g., base stations and / or network switching points) through which one or more components of the application scenario 100 can connect to the network 120 to exchange data and / or information.

[0026] The storage device 130 refers to a device for storing data, instructions, and / or any other information. In some embodiments, the storage device 130 can store data and / or information obtained from the processing device 110, the scanning imaging device 140, and / or the terminal device 150, etc. For example, the storage device 130 can store photon counting data output by a photon counting detector in the scanning imaging device 140. For another example, the storage device 130 can store a scanning image generated by the processing device 110. In some embodiments, the storage device 130 can include one or any combination of mass storage, removable storage, etc.

[0027] The scanning imaging device 140 refers to a device capable of imaging a target part (e.g., a chest, an abdomen, etc.) of a scanning object (e.g., a patient) based on X-rays. For example, the scanning imaging device 140 can be an X-ray flat panel imaging device, a CT (Computed Tomography) device, a PCCT device, a PET (Positron Emission Computed Tomography) device, etc. The scanning imaging device 140 can include an X-ray source (e.g., an X-ray tube), a high-voltage generator, a detector (e.g., a photon counting detector), a gantry, a scanning bed (to carry the scanning object), etc. More details about the structure of the scanning imaging device 140 can be found in Figure 2 and related descriptions.

[0028] Figure 2 is a schematic diagram of an imaging principle of a scanning imaging device according to some embodiments of the present specification.

[0029] As shown in Figure 2 , when the scanning imaging device 140 performs scanning imaging on the scanning object 230, the X-ray source 210 can emit X-rays, which first pass through the collimator 220, then pass through the scanning object 230, and finally reach the photon counting detector 240. The photon counting detector 240 can process the received X-rays to obtain corresponding photon counting data. More details about the structure and principle of the photon counting detector 240 can be found in Figure 3 and related descriptions.

[0030] The terminal device 150 refers to one or more terminals or software used by a user (e.g., a doctor, a researcher, etc. who uses the scanning imaging device 140) of the scanning imaging device 140 or a scanning object. In some embodiments, as shown in Figure 1 , the terminal device 150 can include, but is not limited to, a smart phone 151, a tablet computer 152, a laptop computer 153, a desktop computer 154, etc. In some embodiments, the terminal device 150 can interact with other components in the application scenario 100 through the network 120. For example, the terminal device 150 can obtain a scanning image (e.g., a PCCT image) determined by the processing device 110 and output to the user through the network 120. For another example, the terminal device 150 can control a scanning process of the scanning imaging device 140 (e.g., control the scanning imaging device 140 to start scanning, etc.) through the network 120.

[0031] Figure 3 is an exemplary structural diagram of a photon counting detector according to some embodiments of the present specification.

[0032] As shown in Figure 3As shown, the photon counting detector 240 may include a conversion layer assembly 310, a photoelectric converter 320, a photon counting circuit 330, and a control readout circuit 340. The conversion layer assembly 310 includes a scintillation crystal array 311. The scintillation crystal array 311 is used to convert X-rays into fluorescence. The photoelectric converter 320 is used to convert fluorescence into an electronic signal. The photon counting circuit 330 is used to process the electronic signal to obtain photon counting data. The control readout circuit 340 is used to read out the photon counting data and transmit the photon counting data to an external readout circuit. In some embodiments, the scintillation crystal array 311 is made of a scintillation crystal material, and the emission decay time of the scintillation crystal material is within 25 ns. Further descriptions of the conversion layer assembly 310, the scintillation crystal array 311, the photoelectric converter 320, the photon counting circuit 330, the control readout circuit 340, the external readout circuit, and the scintillation crystal material can be found in the following descriptions.

[0033] The conversion layer assembly 310 refers to the component in the photon counting detector 240 that converts X-rays into fluorescence and transmits the fluorescence to the photoelectric converter 320. For example... Figure 3 As shown, the conversion layer component 310 includes at least a scintillation crystal array 311.

[0034] The scintillation crystal array 311 refers to an array composed of multiple scintillation crystal units used to convert X-rays into fluorescence (by depositing energy through effects such as the photoelectric effect and Compton effect, and then emitting the energy in the form of fluorescence). A scintillation crystal unit refers to a constituent unit of the scintillation crystal array 311, such as... Figure 3 As shown, the scintillation crystal array 311 may include scintillation crystal units 3111, scintillation crystal units 3112, scintillation crystal units 3113, etc. Multiple scintillation crystal units can be arranged in an array to form the scintillation crystal array 311. For example, the aforementioned multiple scintillation crystal units can be fixed in a square array using a frame structure to form the scintillation crystal array 311. Each scintillation crystal unit in the scintillation crystal array 311 can independently receive X-rays and convert them into fluorescence. Each scintillation crystal unit is coupled to a pixel of the photoelectric converter. Specifically, the back side of the scintillation crystal unit (the side of the scintillation crystal unit facing away from the X-ray source) has the same shape and area as the entrance window of the pixel of the photoelectric converter, and the back side of each scintillation crystal unit is aligned with the entrance window of a pixel, so that the fluorescence generated by each scintillation crystal unit can be received by the pixel it is aligned with, thereby improving the accuracy of single-photon energy measurement. More details about the pixels of the photoelectric converter can be found in the following description.

[0035] Figure 4 This is a schematic diagram illustrating the fluorescence conversion principle of a photon counting detector according to some embodiments of this specification.

[0036] Exemplarily, as shown in Figure 4 The X-ray enters the scintillation crystal unit 3111 after passing through the anti-scattering grid 351 of the anti-scattering grid 350; inside the scintillation crystal unit 3111, the photons in the X-ray deposit energy through photoelectric effect or Compton effect, and the energy is then converted into fluorescent photons, which can be emitted and transmitted multiple times inside the scintillation crystal unit 3111, and finally pass out of the scintillation crystal unit 3111 to reach the fluorescent transmission structure 312; then, the fluorescent photons can pass through the fluorescent transmission structure 312 to reach the plurality of photodiodes (for example, the photodiode 401 and the like) in the photoelectric converter 320; the photocathode of each photodiode can emit electrons (photoelectrons) in response to receiving the fluorescence, and the electrons drift and gather under the action of the electric field to form an electronic signal (for example, a current formed by the aforementioned electron convergence). More about the anti-scattering grid 350, the anti-scattering grid 351, the fluorescent transmission structure 312, and the photoelectric converter 320 can be referred to later.

[0037] It can be understood that, Figure 4 The principle of other scintillation crystal units is the same as that of the scintillation crystal unit 3111.

[0038] The shape of the scintillation crystal unit can be a cube, a cuboid, and the like. For example, the shape of the scintillation crystal unit can be a cuboid as shown in Figure 3

[0039] In some embodiments, the aspect ratio of the scintillation crystal array 311 is greater than 2. The aspect ratio refers to the ratio of the thickness of the scintillation crystal unit to the spacing of the adjacent scintillation crystal units. The thickness refers to the length of the scintillation crystal unit along the X-ray incident direction.

[0040] In some embodiments of the present specification, by setting the aspect ratio of the scintillation crystal array 311, it is helpful to limit the intersection of the fluorescence emitted by the scintillation crystal unit between adjacent scintillation crystal units, thereby weakening the optical crosstalk, and it is helpful to the high spatial resolution of the subsequently generated image. Setting the aspect ratio of the scintillation crystal array 311 can also improve the absorption efficiency of the detector for X-rays, while ensuring that the scintillation crystal array 311 works unsaturated at high count rate.

[0041] In some embodiments, the thickness of the scintillation crystal unit is greater than 100 μm. In some embodiments, the thickness of the scintillation crystal unit is greater than 110 μm. In some embodiments, the thickness of the scintillation crystal unit is greater than 120 μm. In some embodiments, the thickness of the scintillation crystal unit can also be other numerical values, which are not limited here.

[0042] ​In some embodiments of the present disclosure, by limiting the thickness of the scintillation crystal units, the X-rays absorbed by the array can release enough photons, which is conducive to improving the imaging quality (e.g., the quality of the PCCT image) based on the subsequent photon counting data.

[0043] In some embodiments, the scintillation crystal array 311 is composed of a scintillation crystal material. For example, each scintillation crystal unit in the scintillation crystal array 311 is a block of scintillation crystal material.

[0044] In some embodiments, the light emission decay time of the scintillation crystal material described above is within 25 ns. The light emission decay time refers to the time taken by the scintillation crystal material from being excited and generating fluorescence to the light emission intensity of the fluorescence decaying to 1 / e (about 36.8%) of the maximum value.

[0045] The smaller the light emission decay time of the scintillation crystal material, the shorter the time the detector can respond to a single photon and generate fluorescence without being affected by photons arriving at a nearby time (e.g., the time difference between the arrival of two photons at the scintillation crystal array is less than the light emission decay time), which is conducive to the detector processing each independent photon incident event faster and more accurately, and then distinguishing and counting single photons. In some embodiments of the present disclosure, by limiting the light emission decay time of the scintillation crystal material to within 25 ns, the photon counting detector can have the ability to work in a high counting rate situation, and achieve accurate energy measurement and counting of single photons.

[0046] In some embodiments, the light yield of the scintillation crystal material is greater than 10000 / MeV (i.e., for every 1 MeV of X-rays received by the scintillation crystal material, the number of photons generated by the scintillation crystal material is greater than 10000).

[0047] For example, the scintillation crystal material that satisfies the light emission decay time within 25 ns and / or the light yield of the scintillation crystal material greater than 10000 / MeV can be used as the scintillation crystal material constituting the scintillation crystal array 310. For example, the scintillation crystal material can include at least one of cerium bromide, lanthanum bromide, barium fluoride, cerium fluoride, lutetium silicate crystal, and yttrium aluminate crystal.

[0048] In some embodiments of the present disclosure, by using the scintillation crystal material made of the above-mentioned material, the light emission decay time of the scintillation crystal material can be within 25 ns, and the cost of the above-mentioned material is relatively low. By reducing the response time of the photon counting detector, the cost of the photon counting detector is also reduced, thereby realizing low-cost single-photon measurement.

[0049] In some embodiments, the scintillation crystal array 311 can be independently generated and then attached (e.g., attached using silicone grease, optical glue, etc.) to the surface of the photoelectric converter 320.

[0050] In some embodiments, the scintillation crystal array 311 is generated directly on the surface of the photoelectric converter 320 (e.g., the surface on which the photocathode of the photodiode in each pixel of the photoelectric converter 320 is located). In some embodiments, the scintillation crystal array 311 is generated directly on the surface of the photoelectric converter 320 in a large area by a crystal growth, a spin-coating process, or a deposition process. More details about the pixel and the photodiode are described later.

[0051] In some embodiments of the present disclosure, the scintillation crystal array is generated directly on the surface of the photoelectric converter in a large area, which can reduce or even eliminate the gaps between the scintillation crystals of the scintillation crystal array manufactured by the "cutting + splicing" process, effectively improve the effective detection area of the detector, and thus improve the detection efficiency and resolution. In addition, the direct growth technology can simplify the production process, so that the thickness of the scintillation crystal array is controllable and dense and uniform, which improves the stability of the scintillation crystal array, and thus realizes the preparation of a large-size and high-quality scintillation crystal array.

[0052] In some embodiments, as shown in FIG. 3B, the conversion layer assembly 310 further includes a fluorescent light transmission structure 312. The fluorescent light transmission structure 312 is used to transmit the fluorescent light converted by the scintillation crystal array 311 to the photoelectric converter 320. Figure 3

[0053] In some embodiments, the fluorescent light transmission structure 312 includes a reflective layer, a light coupling layer, and a microlens layer (not shown in the figure).

[0054] The reflective layer is used to reflect the fluorescent light converted by the scintillation crystal array 311. The material of the reflective layer is a high-reflectivity material. For example, the material of the reflective layer can be barium sulfate, titanium dioxide paint, Teflon film, an evaporated metal layer (e.g., an aluminum film), etc. The reflective layer is located on the front surface and the side surface of each scintillation crystal unit (e.g., the scintillation crystal unit 3111, the scintillation crystal unit 3112, the scintillation crystal unit 3113, etc.) in the scintillation crystal array 311. Among them, the front surface is the surface of the scintillation crystal unit facing the X-ray source, and the side surface is parallel to the incident direction of the X-ray as shown in FIG. 3A. Figure 3

[0055] In some embodiments of the present disclosure, the reflective layer is arranged on the front surface and the side surface of each scintillation crystal unit, so that the scintillation crystal unit can collect as much fluorescent light generated by the scintillation crystal material as possible and guide it to the back surface of the scintillation crystal unit (the surface of the scintillation crystal unit facing away from the X-ray source, i.e., the opposite surface of the front surface of the scintillation crystal unit), and thus as much fluorescent light generated by the scintillation crystal array 311 as possible is transmitted to the photoelectric converter and received by the photoelectric converter, thereby improving the accuracy of the photon energy measurement.

[0056] ​​The light coupling layer is used to fill the gap between the scintillation crystal array 311 and the post-optical element (e.g., the microlens layer or the photoelectric converter 320), so as to inhibit total reflection of the fluorescent light and maximize the transmittance of the fluorescent light. The material of the light coupling layer can be a coupling liquid material with high transparency. It is characterized by high light transmittance. For example, the material of the light coupling layer can be optical glue, optical silicone or transparent epoxy resin, etc.

[0057] The microlens layer is used to converge the fluorescent light passing through the light coupling layer, so as to gather the fluorescent light into the light-sensitive area (i.e., the area where the photodiode is located) of the pixel of the photoelectric converter 320. The material of the microlens layer is a polymer. For example, the material of the microlens layer can be photoresist, etc.

[0058] The light coupling layer and the microlens layer are located between the scintillation crystal array 311 and the photoelectric converter 320, wherein the light coupling layer is located between the scintillation crystal array 311 and the microlens layer, and the microlens layer is located between the light coupling layer and the photoelectric converter 320. Based on this, it can be understood that the fluorescent light generated by the scintillation crystal unit will first penetrate the light coupling layer, and then reach the photoelectric converter 320 after being converged by the microlens layer.

[0059] In some embodiments of the present specification, the light coupling layer arranged between the scintillation crystal array and the photoelectric converter can reduce the loss of the fluorescent light when propagating in the gap between the scintillation crystal array and the photoelectric converter, thereby improving the light transmission efficiency of the gap between the scintillation crystal array and the photoelectric converter. The microlens layer arranged between the scintillation crystal array and the photoelectric converter converges and outputs the photons, which can improve the overall photoelectric transmission efficiency of the photon counting detector.

[0060] The photoelectric converter 320 is used to convert the fluorescent light into an electronic signal.

[0061] The photoelectric converter 320 includes a pixel array. For example, the pixel array can be an active two-dimensional pixel array.

[0062] The pixel array includes a plurality of pixels (also referred to as channels). The pixel refers to a constituent unit of the photoelectric converter 320, and each pixel can independently perform photoelectric conversion. As shown in Figure 3 The pixel array included in the photoelectric converter 320 includes the pixel 321, the pixel 322, the pixel 323, etc.

[0063] In some embodiments, each pixel corresponds to a photodiode (e.g., the photodiode 401). The photodiode includes a photocathode. The photocathode is used to emit photoelectrons in response to receiving the fluorescent light.

[0064] In some embodiments, the photoelectric conversion duration of the photodiode is less than 100 ps. The photoelectric conversion duration refers to the duration that the photodiode takes to convert the received fluorescent light into an electronic signal (i.e., the duration from the arrival of a photon of the fluorescent light at the photodiode to the generation of an electronic signal by the photodiode based on the photon). The photoelectric conversion duration of the photodiode being less than 100 ps, in combination with the light emission decay time of the scintillation crystal material being within 25 ns, enables the fast response of the photon counting detector 240 to the photons, thereby enabling single-photon measurement and counting.

[0065] The smaller the photoelectric conversion duration of the photodiode, the more conducive to the differentiation and counting of single photons by the photon counting detector. Specifically, the smaller the photoelectric conversion duration of the photodiode, the higher the light flux at which the detector can accurately measure (differentiate and count) single photons. In some embodiments of the present disclosure, by limiting the photoelectric conversion duration of the photodiode to be within 100 ps, in combination with the light emission decay time of the scintillation crystal material being within 25 ns, the accuracy of the single-photon measurement and counting by the photon counting detector can be further improved.

[0066] In some embodiments, the operating voltage of the photodiode can be configured. The gain mode of the photodiode is related to the value of the operating voltage of the photodiode, i.e., the gain mode of the photodiode can change in response to the change of the operating voltage. Exemplarily, the operating voltage can be applied to the P-N junction of the photodiode, and the operating voltage can be a reverse bias voltage.

[0067] In some embodiments, the gain mode of the photodiode includes a photoconductive mode and a linear amplification mode. When the operating voltage of the photodiode is configured to be zero bias or low reverse bias (e.g., less than 50 V), the operating mode of the photodiode is the photoconductive mode; when the operating voltage of the photodiode is configured to be high reverse bias (e.g., greater than 50 V), the gain mode of the photodiode is the linear amplification mode.

[0068] In some embodiments, when the gain mode of the photodiode is the photoconductive mode, the gain value of the photodiode is 1.

[0069] The gain value refers to the ratio of the number of photoelectrons emitted by the photocathode of the photodiode to the number of photons of the fluorescent light received by the photodiode. For example, if the gain value of the photodiode is 1, it means that for every 1 photon of the fluorescent light received by the photodiode, 1 photoelectron can be emitted by the photocathode of the photodiode. For another example, if the gain value of the photodiode is 5, it means that for every 1 photon of the fluorescent light received by the photodiode, 5 photoelectrons can be emitted by the photocathode of the photodiode.

[0070] In some embodiments, the gain value of the photodiode is greater than 1 when the gain mode of the photodiode is the linear amplification mode. For example, the gain value of the photodiode can be any integer value in the range of [2, 1000] when the gain mode of the photodiode is the linear amplification mode.

[0071] A conventional photoelectric converter (e.g., a solid-state photomultiplier (SSPM) or a silicon photomultiplier (SiPM)) has a structure of an array of multiple single-photon avalanche diodes (SPADs) connected in parallel. The gain value of a single-photon avalanche diode is a large fixed value (e.g., 10 6 It can be understood that the number of electrons (photoelectrons) generated by a single-photon avalanche diode is proportional to the number of incident photons (the proportional coefficient is the gain value), and the intensity (e.g., voltage value) of the electronic signal generated by a single-photon avalanche diode is proportional to the number of electrons, and thus the intensity of the electronic signal generated by a single-photon avalanche diode is proportional to the number of incident photons, and the signal output amplitude (e.g., voltage value) of a conventional photoelectric converter is proportional to the number of single-photon avalanche diodes triggered by photons. Based on this, the conventional photoelectric converter can inversely deduce the total number of incident photons based on the signal output amplitude through the above proportional relationship, thereby realizing linear counting of photons. However, due to the large gain value of a single-photon avalanche diode, problems such as high power consumption and serious heat generation may occur, and there may be a problem of micro-light link triggering, which leads to poor anti-optical crosstalk capability of the conventional photoelectric converter. In addition, the structure of a solid-state photomultiplier or a silicon photomultiplier is complex, the manufacturing process is difficult, and the manufacturing cost is high. In some embodiments of the present specification, different voltages are configured for photodiodes, so that the gain modes (or gain values) of the photodiodes are different, thereby avoiding the photodiodes always working in a “high gain value” state, greatly reducing the power consumption, heat generation, and manufacturing difficulty of the photoelectric converter, and improving the anti-optical crosstalk capability.

[0072] In some embodiments, the fill factor of each pixel is greater than 50%. The fill factor refers to the ratio of the area of the light-sensitive surface of the photodiode to the area of the pixel. The pixel area refers to the area of the face of the pixel directly opposite the scintillation crystal array 311.

[0073] Each pixel in the photoelectric converter 320 is coupled with each scintillation crystal unit in the scintillation crystal array 311 for receiving the fluorescent light generated by the coupled scintillation crystal unit and converting the fluorescent light into an electronic signal. In some embodiments, the pixels of the photoelectric converter 320 can be coupled one-to-one with the scintillation crystal units of the scintillation crystal array 311. For example, as shown in FIG. 3, the photoelectric converter 320 has a pixel array 321, and the pixel array 321 has a one-to-one correspondence with the scintillation crystal array 311.Figure 3 As shown, pixel 321 is coupled with scintillation crystal unit 3111, pixel 322 is coupled with scintillation crystal unit 3112, pixel 323 is coupled with scintillation crystal unit 3113, and so on. At this time, each pixel in photoelectric converter 320 is used to receive the fluorescence generated by the single scintillation crystal unit it is coupled with, and convert the fluorescence into an electronic signal. Based on this, in some embodiments, the area of a pixel can be equal to the area of the back surface of the scintillation crystal unit corresponding to the pixel. In some embodiments, each pixel of photoelectric converter 320 is coupled with multiple scintillation crystal units in scintillation crystal array 311, and the pixel can receive the fluorescence generated by the multiple scintillation crystal units it is coupled with, and convert the fluorescence into an electronic signal.

[0074] In some embodiments, the equivalent electronic noise value of a pixel in photoelectric converter 320 is positively correlated with the area of the pixel. That is, the larger the area of a pixel, the greater the equivalent electronic noise value of the pixel.

[0075] The equivalent electronic noise value of a pixel refers to the noise value obtained by equivalent processing of the noise caused by the electronic elements (e.g., photodiodes) and internal circuit structures contained in the pixel.

[0076] The area of a pixel refers to the area of the part of the pixel that can receive fluorescence.

[0077] Exemplarily, the relationship between the equivalent electronic noise value of a pixel and the area of the pixel is shown in the following formula (1):

[0078] wherein, the equivalent electronic noise value of the pixel, the area of the pixel, the upper limit of the equivalent electronic noise value of the pixel, the value of a can be preset.

[0079] In some embodiments, the upper limit of the equivalent electronic noise value of a pixel is 2000 electrons per square millimeter. Exemplarily, when the area of a pixel is 0.25 mm 2 , the upper limit of the equivalent electronic noise value of the pixel is 500 electrons, that is, when the area of a pixel is 0.25 mm 2 , the equivalent electronic noise value of the pixel does not exceed 500 electrons.

[0080] In some embodiments of the present specification, by limiting the equivalent electronic noise value of a pixel, the noise of the pixel is prevented from being too large, and the accuracy of the photon counting detector is ensured; by using a photoelectric converter with a very small photoelectric conversion time (e.g., less than 100 ps), the response speed of the photon counting detector is accelerated, so that the photon counting detector can meet the single photon counting requirement in a clinical application scenario.

[0081] In some embodiments, each pixel in the photoelectric converter 320 is connected with a signal processing circuit respectively. More details about the signal processing circuit can be found in the relevant description below.

[0082] The photon counting circuit 330 is configured to process the electronic signal generated by the photoelectric converter 320 to obtain photon counting data. The processing includes pre-processing, amplitude discrimination and photon counting, wherein the pre-processing includes amplification, filtering. The photon counting data includes a plurality of photon counting sub-data. More details about the photon counting sub-data can be found in the relevant description below.

[0083] In some embodiments, the photoelectric converter 320 is manufactured by an integrated process. For example, a photolithography technology can be used to photoetch each pixel corresponding to each scintillation crystal unit on a wafer substrate, thereby obtaining the pixel array in the photoelectric converter 320. Compared with separately installing a corresponding pixel for each scintillation crystal unit, directly photoetching each pixel corresponding to each scintillation crystal unit on the wafer substrate reduces the difficulty of corresponding installation of the scintillation crystal unit and the pixel, and avoids mutual interference of fluorescence signals in different pixels, thereby improving the accuracy of photon detection.

[0084] In some embodiments, the photoelectric converter and the photon counting circuit are manufactured by an integrated process. More details about the foregoing embodiments can be found in the relevant description below.

[0085] Figure 5 is a circuit structure schematic diagram of the photon counting circuit according to some embodiments of the present specification.

[0086] The photon counting circuit 330 includes a plurality of signal processing circuits, a plurality of threshold comparator groups and a plurality of counter groups. For example Figure 5 As shown, the plurality of signal processing circuits includes a signal processing circuit 511, a signal processing circuit 512 and a signal processing circuit 513, etc., the plurality of threshold comparator groups includes a threshold comparator group 521, a threshold comparator group 522 and a threshold comparator group 523, etc., and the plurality of counter groups includes a counter group 531, a counter group 532 and a counter group 533, etc.

[0087] The plurality of signal processing circuits are respectively connected with the plurality of pixels in the pixel array to respectively receive the electronic signal output by each pixel. For example Figure 5 As shown, the signal processing circuit 511 is connected with the pixel 321 to receive the electronic signal output by the pixel 321; the signal processing circuit 512 is connected with the pixel 322 to receive the electronic signal output by the pixel 322; and the signal processing circuit 513 is connected with the pixel 323 to receive the electronic signal output by the pixel 323.

[0088] The signal processing circuit is configured to pre-process the electronic signal output by the photoelectric converter. The pre-processing includes amplification and filtering.

[0089] In some embodiments, the signal processing circuit includes a preamplifier and a filter, wherein the preamplifier is configured to amplify the electronic signal, and the filter is configured to filter the electronic signal.

[0090] Figure 6 is a circuit structure schematic diagram of a photon counting circuit according to another embodiment of the present specification.

[0091] Exemplarily, as shown in Figure 6 the signal processing circuit 511 includes a preamplifier 611 and a filter 612, wherein the preamplifier 611 is configured to amplify the electronic signal, and the filter 612 is configured to filter the electronic signal.

[0092] In some embodiments, as shown in Figure 6 the preamplifier 611 is provided with a dynamic reset function. The dynamic reset refers to that the preamplifier 611 releases the accumulated charge after amplifying one electronic signal, so as to be ready for processing the next electronic signal.

[0093] In some embodiments, as shown in Figure 6 the preamplifier 611 is connected in parallel with a capacitor C, and the capacitor C is configured to prevent circuit oscillation and improve the stability of the signal processing circuit 511.

[0094] The signal processing circuit 511 further includes a leakage compensator 613, and the leakage compensator 613 is configured to compensate for the lost charge of the capacitor C. In this way, the amount of charge of the capacitor C is stabilized.

[0095] In some embodiments, as shown in Figure 6 the photon counting circuit 330 further includes a time trigger 620. The time trigger 620 is configured to record the time when the filtering of the electronic signal is completed (i.e., the time when the pre-processing is completed), the time when the preamplifier 611 is reset, and the like.

[0096] The plurality of groups of threshold value comparators are respectively connected with the plurality of signal processing circuits, so as to respectively receive the pre-processed electronic signals output by the plurality of signal processing circuits.

[0097] Each of the plurality of groups of threshold value comparators includes a plurality of threshold value comparators. Each threshold value comparator is configured to perform amplitude discrimination. For example Figure 5As shown, threshold comparator group 521 includes threshold comparator 5211, threshold comparator 5212 and threshold comparator 5213, threshold comparator group 522 includes threshold comparator 5221, threshold comparator 5222 and threshold comparator 5223, and threshold comparator group 523 includes threshold comparator 5231, threshold comparator 5232 and threshold comparator 5233.

[0098] The threshold comparator is used to perform amplitude discrimination, including: determining whether the amplitude (or level value) of the preprocessed electronic signal is higher than the amplitude threshold preset by the threshold comparator; and generating a digital pulse in response to the amplitude of the preprocessed electronic signal being higher than the amplitude threshold of the threshold comparator.

[0099] The amplitude threshold refers to a preset voltage value for the threshold comparator. If the voltage value of the electronic signal received by the threshold comparator is higher than the preset amplitude threshold, the threshold comparator generates a digital pulse; if the voltage value of the electronic signal received by the threshold comparator is lower than the preset amplitude threshold, the threshold comparator does not generate a digital pulse.

[0100] The amplitude thresholds of the various threshold comparators within the same threshold comparator group are different. For example, such as... Figure 6 As shown, the threshold comparator group 521 includes threshold comparator 5211, threshold comparator 5212, and threshold comparator 5213, wherein the amplitude thresholds of threshold comparators 5211, 5212, and 5213 increase sequentially. For example, the amplitude threshold V of threshold comparator 5211... ref1 The amplitude threshold V of the threshold comparator 5212 is 1mV. ref2 The amplitude threshold V of the threshold comparator 5213 is 2mV. ref3 It is 3mV.

[0101] The multiple counters of the photon counting circuit 330 are connected to multiple threshold comparators to count the digital pulses generated by each threshold comparator.

[0102] Each of the multiple sets of counters includes multiple counters, wherein the multiple counters in each set are connected to multiple threshold comparators in a set of threshold comparators connected to that set of counters, and each counter is used to perform photon counting. For example Figure 5As shown, the counter group 531 includes the counter 5311, the counter 5312 and the counter 5313, the counter group 532 includes the counter 5321, the counter 5322 and the counter 5323, and the counter group 533 includes the counter 5331, the counter 5332 and the counter 5333. Among them, the counter 5311, the counter 5312 and the counter 5313 in the counter group 531 are connected with the threshold comparator 5211, the threshold comparator 5212 and the threshold comparator 5213 in the threshold comparator group 521 respectively, the counter 5321, the counter 5322 and the counter 5323 in the counter group 532 are connected with the threshold comparator 5221, the threshold comparator 5222 and the threshold comparator 5223 in the threshold comparator group 522 respectively, and the counter 5331, the counter 5332 and the counter 5333 in the counter group 533 are connected with the threshold comparator 5231, the threshold comparator 5232 and the threshold comparator 5233 in the threshold comparator group 523 respectively.

[0103] The counter is used to perform photon counting, including counting the digital pulses generated by the connected threshold comparators to determine the photon counting sub-data.

[0104] The photon counting sub-data refers to the counting data of the digital pulses generated by the threshold comparator group connected with each group of counters. In combination with the foregoing, each pixel corresponds to a group of threshold comparators and a group of counters respectively, and it can be understood that the photon counting sub-data obtained by each group of counters can also be regarded as the photon counting sub-data corresponding to the pixel corresponding to the group of counters.

[0105] Exemplarily, referring to Figures 3 to 6 , still taking the amplitude threshold V ref1 as 1 mV, the amplitude threshold V ref2 as 2 mV, and the amplitude threshold V ref3 as 3 mV as an example, it is assumed that the number of the preprocessed electronic signals received by the threshold comparator group 521 is 10, and the voltage values of the 10 preprocessed electronic signals are 1.5 mV, 2.2 mV, 0.1 mV, 3.3 mV, 4 mV, 3.2 mV, 1.7 mV, 2.8 mV, 4.3 mV and 1.9 mV respectively. Among them, there is 1 voltage value less than V ref1 , there are 3 voltage values greater than V ref1 and less than V ref2 , there are 2 voltage values greater than V ref2 and less than V ref3 , and there are 4 voltage values greater than V ref3Therefore, the counting data of the counter 5311 is 3+2+4=9, the counting data of the counter 5312 is 2+4=6, and the counting data of the counter 5313 is 4. Based on this, the photon counting sub-data of the counter group 531 is 9, representing that the pixel 321 corresponding to the counter group 531 receives a total of 9 photons.

[0106] In some embodiments, the photon counting sub-data of each counter group can also distinguish the counting data of each counter in the counter group. For example, still taking the counting data of each counter in the counter group 531 in the previous example, the photon counting sub-data of the counter group 531 is (3, 2, 4), representing that among all the preprocessed electronic signals received by the threshold comparator group 521 corresponding to the counter group 531, the voltage values of 3 electronic signals are greater than V ref1 and less than V ref2 , the voltage values of 2 electronic signals are greater than V ref2 and less than V ref3 , and the voltage values of 4 electronic signals are greater than V ref3 .

[0107] It can be understood that when the energy of the X-ray received by the scintillation crystal unit is higher, the energy of the fluorescent photons converted by the scintillation crystal unit is also higher, and the voltage value of the electronic signal converted by the pixel of the photoelectric converter is also higher, and the preprocessed electronic signal also has a larger voltage value. A larger voltage value can cause the amplitude threshold value to be larger, and the threshold comparator to generate a digital pulse, so that the counter corresponding to the threshold comparator with a larger amplitude threshold value generates a count. Based on the foregoing principle, the photon counting sub-data can represent the number of photons in different energy ranges (referred to as energy windows). For example, it is assumed that the amplitude threshold value V ref1 corresponds to a photon energy E1 (for example, 1.5 keV), the amplitude threshold value V ref2 corresponds to a photon energy E2 (for example, 3 keV), and the amplitude threshold value V ref3Corresponding photon energy is E3 (for example, 4.5 keV), then the photon counting sub-data (3, 2, 4) described in the foregoing example represents that, among the 9 photons received by the pixel 321, 3 photons have energy greater than E1 and less than E2, 2 photons have energy greater than E2 and less than E3, and 4 photons have energy greater than E3. That is, the photon counting sub-data includes counting results of photons in multiple different energy intervals (energy windows). Different material components in the scanned object (for example, the scanned object 230) have different degrees of weakening to X-ray energy, so that X-rays reaching the scintillation crystal unit have different energies, and the scintillation crystal unit generates fluorescent photons with different energies. Through the above energy distinguishing manner, the scanning imaging device (for example, the scanning imaging device 140) with the energy distinguishing function can identify different material components (for example, fat, bone, water, etc.) in the scanned object (for example, the scanned object 230). For example, the energy data corresponding to the photon counting data can be analyzed by a material decomposition algorithm to identify the material components in the scanned object.

[0108] In some embodiments, the photon counting data includes multiple photon counting sub-data. For example, the photon counting data includes photon counting sub-data of all the counter groups. For example, it is assumed that the photon counting circuit includes 4 counter groups in total, and the photon counting sub-data of each counter group is 8, 11, 14, and 9 respectively, then the photon counting data is (8, 11, 14, 9). For another example, it is assumed that the photon counting circuit includes 4 counter groups in total, and the photon counting sub-data of each counter group is (3, 2, 4), (5, 5, 2), (1, 2, 3), and (4, 1, 3) respectively, then the photon counting data is [(3, 2, 4), (5, 5, 2), (1, 2, 3), (4, 1, 3)].

[0109] In some embodiments, the photon counting circuit 330 is manufactured through an integrated process.

[0110] In some embodiments, the photoelectric converter 320 and the photon counting circuit 330 are manufactured through an integrated process. For example, the photoelectric converter 320 can be obtained by photolithography technology, that is, each pixel corresponding to each scintillation crystal unit is photolithographed on a wafer substrate, and then each signal processing circuit, threshold comparator, and counter corresponding to each pixel is photolithographed on the wafer substrate. In some embodiments, for example, the integrated process includes a complementary metal oxide semiconductor (CMOS) process. The photoelectric converter and the photon counting circuit manufactured through the integrated process reduce the structural complexity and packaging process difficulty of the photon counting detector, and improve the reliability of the photon counting detector.

[0111] Figure 7 Figure 1 is a structural schematic diagram of a photoelectric converter and a photon counting circuit integrated according to some embodiments of the present specification.

[0112] As shown in Figure 7 , the photoelectric converter and the photon counting circuit are integrated on the same wafer substrate through an integration process, and the photoelectric converter and the photon counting circuit form an integral whole, achieving a very high integration of the photoelectric converter and the photon counting circuit.

[0113] In some embodiments of the present specification, the photoelectric converter and the photon counting circuit are integrated directly on a wafer, eliminating the flip-chip process in the traditional manufacturing method, which not only achieves a high integration of the photoelectric converter and the photon counting circuit, reduces the parasitic capacitance of the photon counting detector, weakens signal crosstalk, but also enables large-scale and high-consistency production of the photoelectric converter and the photon counting circuit, while avoiding the complex packaging process caused by the use of splicing technology (e.g., flip-chip), significantly reducing the processing difficulty and cost of a single detection unit (e.g., a single pixel, a single signal processing circuit).

[0114] In some embodiments, the output interface (not shown in the figure) of the photon counting circuit 330 is in communication with the packaging electrode in the control readout circuit 340 to realize signal transmission between the photon counting circuit 330 and the control readout circuit 340. In some embodiments, the above-mentioned output interface is made through a through-silicon via (TSV) process. For more information about the control readout circuit 340, please refer to the relevant description later.

[0115] In some embodiments of the present specification, the output interface of the photon counting circuit is made through a through-silicon via process instead of a front-side wire bonding process, which realizes vertical close interconnection between the photon counting circuit and the control readout circuit, and lays a foundation for obtaining a large-area detector with almost no dead zone; in addition, the through-silicon via process creates a shorter electrical connection path for the photon counting circuit and the control readout circuit, which helps to reduce delay, power consumption and crosstalk in signal transmission, and improve signal integrity and transmission speed.

[0116] The control readout circuit 340 is used to read out photon counting data and transmit the photon counting data to an external readout circuit. The external readout circuit is the internal circuit of a device or system that needs to use photon counting data, for example, the internal circuit of the processing device 110 or the scanning imaging device 140.

[0117] In some embodiments, the control readout circuit 340 can read out the photon counting sub-data corresponding to each pixel respectively, and determine the photon counting data based on all the photon counting sub-data. The method for determining the photon counting data can be referred to the description above.

[0118] In some embodiments, the pixel array includes a plurality of pixel blocks, wherein each pixel block includes a plurality of pixels. For example, taking a 9x9 pixel array (a pixel array formed by 9 rows and 9 columns of pixels) as an example, the pixel array can include 9 3x3 pixel blocks, each of which is composed of 3 rows and 3 columns of pixels.

[0119] In some embodiments, the control readout circuit 340 can read out the photon counting sub-data corresponding to each pixel respectively, and determine the photon counting data based on the photon counting sub-data. For example, the photon counting data includes the photon counting sub-data. The photon counting sub-data corresponding to each pixel includes the photon counting sub-data of each pixel in the pixel block. For example, taking a pixel block composed of 3 rows and 3 columns of pixels as an example, the pixel block includes 9 pixels, and assuming that the photon counting sub-data of the aforementioned 9 counter groups are 8, 11, 14, 9, 7, 11, 10, 16, and 11, respectively, then the photon counting sub-data corresponding to the pixel block is (8, 11, 14, 9, 7, 11, 10, 16, 11).

[0120] In some embodiments of the present specification, reading out the photon counting sub-data corresponding to each pixel respectively can ensure the accuracy of the photon counting data, and reading out the photon counting sub-data corresponding to each pixel block respectively can improve the reading efficiency of the photon counting data while ensuring the accuracy of the photon counting data.

[0121] In some embodiments, as shown in FIG. 3B, the photon counting detector 240 further includes an anti-scattering grille (ASG) 350, which is arranged on the side of the scintillation crystal array 311 opposite the X-ray source, so that the X-rays first pass through the anti-scattering grille 350 and then enter the scintillation crystal array 311. Figure 3

[0122] The anti-scattering grille 350 is a grating structure capable of absorbing scattered X-rays passing through a scanning object (for example, a human body), which can prevent scattered X-rays from entering the scintillation crystal array 311.

[0123] The anti-scattering grille 350 is made of an anti-scattering material, which refers to a material capable of shielding or absorbing X-rays. For example, the material of the anti-scattering grille 350 can be at least one of lead, molybdenum, and tungsten, or an alloy containing at least one of lead, molybdenum, and tungsten.

[0124] The anti-scattering grille 350 includes a grating wall and a plurality of transmission holes, wherein the plurality of transmission holes are surrounded by the grating wall. For example Figure 3 ​As shown, the anti-scattering grating 350 includes grating walls 3501 and a plurality of transmission holes, where the plurality of transmission holes includes transmission hole 351, transmission hole 352, and the like.

[0125] In some embodiments, among the plurality of transmission holes of the anti-scattering grating 350, the pitch of adjacent transmission holes (for example, the pitch of transmission hole 351 and transmission hole 352) is equal to the pitch of adjacent pixels in the pixel array of the photoelectric converter.

[0126] In some embodiments, the thickness of the grating wall 3501 is less than 150 μm. In some embodiments, the thickness of the grating wall 3501 is less than 140 μm. In some embodiments, the thickness of the grating wall 3501 is less than 130 μm. In some embodiments, the thickness of the grating wall 3501 can also be other values, which are not limited herein.

[0127] Each transmission hole of the anti-scattering grating 350 corresponds to a respective scintillation crystal unit in the scintillation crystal array 311, so that the X-ray passing through the transmission hole can only reach the corresponding scintillation crystal unit. For example Figure 3 As shown, the anti-scattering grating 351 corresponds to the scintillation crystal unit 3111, and the X-ray passing through the transmission hole 351 can only reach the scintillation crystal unit 3111; the anti-scattering grating 352 corresponds to the scintillation crystal unit 3112, and the X-ray passing through the transmission hole 352 can only reach the scintillation crystal unit 3112.

[0128] In some embodiments of the present specification, by setting the anti-scattering grating, the scattering of X-rays can be effectively prevented, and the photon counting error caused by the scattering of X-rays can be eliminated. By limiting the thickness of the grating wall to a small value (for example, less than 150 μm), the effect of preventing the scattering of X-rays can be achieved while reducing the shielding area of the grating wall to the scintillation crystal array, thereby reducing the dead zone of the scintillation crystal array.

[0129] In some embodiments of the present specification, the photon counting detector with the above structure solves the problems of complex process, high cost, high image noise, unstable operation and easy polarization of the mainstream photon counting detector, and has cost advantage and process advantage.

[0130] In some embodiments of the present specification, the scintillation crystal material with a specific light-emitting decay time (for example, within 25 ns) and the photoelectric converter with a specific photoelectric conversion time (for example, less than 100 ps) are combined to achieve accurate observation of X-rays and single-photon counting; the photon counting detector described in some embodiments of the present specification uses the above-mentioned photon counting circuit and control readout circuit, so that the conversion mode of the above-mentioned photon counting detector for the detection signal (for example, an electronic signal converted based on X-rays) is different from that of the traditional detector, thereby improving the sensitivity of the detector and the accuracy of the detection signal, while having the characteristics of low noise, simple process, large-area lithography, high spatial resolution, fast response speed, etc., and can meet the single-photon counting requirements in the clinical application scenario.

[0131] The above has described the basic concept, and it is obvious that the above detailed disclosure is only used as an example and does not limit the present specification. Although it is not explicitly stated here, those skilled in the art can make various modifications, improvements and corrections to the present specification. Such modifications, improvements and corrections are suggested in the present specification, so such modifications, improvements and corrections still belong to the spirit and scope of the exemplary embodiments of the present specification.

[0132] At the same time, specific words are used in the present specification to describe the embodiments of the present specification. As "one embodiment", "an embodiment", and / or "some embodiments" means a certain feature, structure or characteristic related to at least one embodiment of the present specification. Therefore, it should be emphasized and noted that "an embodiment" or "one embodiment" or "one alternative embodiment" mentioned in different places in the present specification does not necessarily refer to the same embodiment. In addition, some features, structures or characteristics in one or more embodiments of the present specification can be properly combined.

[0133] In addition, unless the claim explicitly states otherwise, the order of the processing elements and sequences described in the present specification, the use of numerals and letters, or the use of other names, is not intended to limit the order of the processes and methods of the present specification. Although some currently considered useful embodiments of the invention are discussed in the above disclosure through various examples, it should be understood that such details are only for the purpose of illustration, and the additional claims are not limited to the disclosed embodiments, on the contrary, the claims are intended to cover all modifications and equivalent combinations that meet the spirit and scope of the embodiments of the present specification. For example, although the system components described above can be realized by hardware devices, they can also be realized by only software solutions, such as installing the described system on existing servers or mobile devices.

[0134] For simplicity and to facilitate understanding of one or more embodiments, a description of an embodiment sometimes refers to a plurality of features in a single embodiment, drawing, or description of an embodiment. However, this method of disclosure is not to be interpreted as meaning that the claimed embodiment requires more features than are explicitly recited in the claims. In fact, claims that do not specifically claim a combination of features are intended to cover the various possible combinations of features as would be understood by a person of ordinary skill in the art.

[0135] Some embodiments use numerical values to describe components, quantities of attributes. It should be understood that such numerical values used in the description of embodiments are, in some examples, modified by the adjectives "about," "approximately," or "substantially." Unless otherwise stated, "about," "approximately," or "substantially" indicate that the described numerical value allows for a variation of ±20%. Accordingly, numerical values used in the specification and claims of some embodiments are approximations that can vary depending on the desired characteristics of the individual embodiments. In some embodiments, numerical values used in the specification and claims are approximations that can vary depending on the desired characteristics of the individual embodiments. In some embodiments, numerical values should be considered in the context of the number of significant digits used in the number and the accepted bits of precision of the number. Although the numerical ranges and parameters setting forth the broadest scope of some embodiments of the specification are approximations, the numerical values set forth in the specific examples are reported as precisely as reasonably possible. The application is not limited to the specific numerical values set forth in the examples.

[0136] Each patent, patent application, patent publication, and other material cited in this specification is hereby incorporated by reference in its entirety. In the event of inconsistencies between the disclosure of this specification and the materials incorporated by reference, the disclosure of this specification shall prevail. In the event of inconsistencies between the disclosure of this specification and the claims, the claims shall prevail. In the event of inconsistencies between the disclosure of this specification and the materials incorporated by reference (whether attached hereto or subsequently added), the disclosure of this specification shall prevail. In the event of inconsistencies between the disclosure of this specification and the description, definitions, and / or terminology used in the materials incorporated by reference, the description, definitions, and / or terminology used in this specification shall prevail.

[0137] Finally, it should be understood that the embodiments described herein are merely exemplary of the principles of the embodiments described herein. Other variations having essentially the same structure and function but different values for components, and / or different arrangements of the components can also be utilized. Accordingly, the embodiments described herein are not to be considered as limited to the examples described herein.

Claims

1. A photon counting detector, characterized in that, include: Scintillation crystal arrays are used to convert X-rays into fluorescence; A photoelectric converter for converting the fluorescence into an electronic signal; A photon counting circuit is used to process the electronic signal to obtain photon counting data; A control readout circuit is used to read out the photon counting data and transmit the photon counting data to an external readout circuit; The scintillation crystal array is made of scintillation crystal material, and the light emission decay time of the scintillation crystal material is within 25 ns.

2. The photon counting detector as described in claim 1, characterized in that, The scintillation crystal array includes multiple scintillation crystal units, and the aspect ratio of the scintillation crystal array is greater than 2. The aspect ratio is the ratio of the thickness of the scintillation crystal unit to the spacing between adjacent scintillation crystal units.

3. The photon counting detector as described in claim 1, characterized in that, The photoelectric converter includes a pixel array, which comprises multiple pixels, wherein... Each pixel includes a photodiode, the operating voltage of which can be configured, and the gain mode of which is related to the value of the operating voltage of the photodiode; The fill factor of each of the plurality of pixels is greater than 50%, and the fill factor is the ratio of the area of ​​the photosensitive surface of the photodiode to the area of ​​the pixel.

4. The photon counting detector as described in claim 3, characterized in that, The processing includes amplitude discrimination and photon counting. The photon counting data includes multiple photon count sub-data. The photon counting circuit includes multiple signal processing circuits, multiple sets of threshold comparators, and multiple sets of counters. The plurality of signal processing circuits are respectively connected to the plurality of pixels in the pixel array, and the signal processing circuits include a preamplifier and a filter; The plurality of threshold comparators are respectively connected to the plurality of signal processing circuits. Each of the plurality of threshold comparators includes a plurality of threshold comparators, wherein each threshold comparator is used to perform the amplitude discrimination, including: Determine whether the amplitude of the electronic signal is higher than the preset amplitude threshold of the threshold comparator; A digital pulse is generated in response to the amplitude of the electronic signal being higher than the amplitude threshold of the threshold comparator; The plurality of counters are respectively connected to the plurality of threshold comparators. Each of the plurality of counters includes a plurality of counters, wherein the plurality of counters in each group are respectively connected to the plurality of threshold comparators in the group of threshold comparators connected to that group of counters. Each counter is used to perform the photon counting, including: The digital pulses generated by the connected threshold comparator are counted to determine the photon count sub-data.

5. The photon counting detector as described in claim 4, characterized in that, The pixel array includes multiple pixel blocks, wherein each pixel block includes multiple pixels; The readout control circuit reads out the photon count sub-data corresponding to each pixel, or reads out the photon count partition data corresponding to each pixel block.

6. The photon counting detector as described in claim 2, characterized in that, It also includes a fluorescence transmission structure, which includes a reflective layer, an optical coupling layer, and a microlens layer. The reflective layer is located on the front surface and the side surface of each scintillation crystal unit. The front surface is the side of the scintillation crystal unit facing the X-ray source, and the side surface is parallel to the incident direction of the X-rays. The optical coupling layer and the microlens layer are located between the scintillation crystal array and the photoelectric converter.

7. The photon counting detector as described in claim 3, characterized in that, It also includes an anti-scattering grating, which is located on the side of the scintillation crystal array facing the X-ray source. The anti-scattering grating includes a grating wall and a plurality of transmission holes, wherein the spacing between adjacent transmission holes is equal to the spacing between adjacent pixels in the pixel array; the thickness of the grating wall is less than 150 μm.

8. The photon counting detector as claimed in claim 1, characterized in that, The scintillation crystal array is generated directly on the surface of the photoelectric converter, and the photoelectric converter and the photon counting circuit are manufactured using an integrated process.

9. The photon counting detector as claimed in claim 1, characterized in that, The output interface of the photon counting circuit is connected to the packaged electrode in the control readout circuit to realize signal transmission between the photon counting circuit and the control readout circuit. The output interface is fabricated using through-silicon via (TSV) technology.

10. A scanning imaging device, characterized in that, include: X-ray source; as well as The photon counting detector as described in any one of claims 1 to 9.

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