Systems and methods for computed tomography - Patents.com

Abstract

To provide systems and methods for increasing the quality of computed tomography (CT) images.SOLUTION: In one embodiment, a computed tomography (CT) detector system comprises: a layer of energy integrating detectors (EIDs) arranged below a layer of photon counting (PC) sensors with respect to an incoming x-ray, where the number of the PC sensors exceeds the number of the EIDs; and an image processing unit configured to correct PC data using EID data, and denoise and increase the resolution of an image reconstructed from EID data and PC data using a deep learning convolutional neural network (CNN) trained on multiple pairs of images. Each pair of images includes a target image reconstructed from a first signal from the layer of PC sensors, and an input image reconstructed from a second signal from the layer of EIDs.SELECTED DRAWING: Figure 5A

Description

[Technical field]

[0001] FIELD OF THE DISCLOSURE

[0002] Embodiments of the subject matter disclosed herein relate to medical imaging, and more particularly, to improving the quality of images reconstructed using computed tomography imaging systems. [Background technology]

[0002] In a computed tomography (CT) imaging system, a beam of electrons generated by a cathode is directed at a target inside an x-ray tube. The electrons strike the target, producing a fan or cone beam of x-rays that are then directed at an object being examined (such as a patient). After being attenuated by the object, the x-rays strike an array of radiation detectors, producing an image.

[0003] The quality of CT images can be improved by using photon-counting CT (PCCT). In PCCT, the radiation detector is a photon-counting detector, which counts photons to obtain spectral information. PCCT uses a direct conversion detector, which has various advantages over conventional energy-integrating detector (EID) CT systems that use indirect conversion detectors. However, in PCCT systems, due to the limited performance of the photon-counting detector, photon pile-up can occur at high input count rates, resulting in poor image quality at high X-ray flux rates. To reduce the effect of this pile-up, it is possible to introduce smaller detector pixels. However, reducing the size of the detector pixel increases the number of channels used, which may increase both power consumption and data size. Smaller pixels result in more charge sharing between channels because the pixels are closer together, and the gap between detector modules may be comparable to the detector pixel size. Furthermore, there is a risk of more pixels with poor performance. Summary of the Invention

[0004] The present disclosure addresses at least one or more of the above identified problems with a CT detector system including a layer of energy integrating detectors (EIDs) disposed on an opposite side of a layer of photon counting (PC) sensors with respect to incident x-rays, where the number of PC sensors exceeds the number of EID detectors. The CT detector system may further include an image processing unit configured to correct the PC data using the EID data and denoise an image reconstructed from the EID data and the PC data using a deep learning convolutional neural network (CNN) trained on a plurality of pairs of images to improve image resolution. Each pair of images includes a target image reconstructed from a first signal from the layer of PC sensors and an input image reconstructed from a second signal from the layer of EID detectors, where the EID data and the PC data are obtained from the EID data and the PC data acquired simultaneously from the same patient x-ray path. By using a multi-layer CT detector including a layer of EID detectors (e.g., a two-dimensional EID detector array) and a layer of PC sensors (e.g., a two-dimensional PC sensor array), the advantages of both PC and EID detectors can be utilized to achieve higher image quality than that achieved by either detector type alone for various clinical CT applications. By placing the layer of PC sensors between the layer of EID detectors and the x-ray source, the size of the PC sensors can be reduced compared to conventional PCCT detector systems, which reduces the effects of pile-up and increases the resolution and reduces the amount of noise in images reconstructed from both PC and EID data compared to images reconstructed from either PC or EID data alone.

[0005] Furthermore, by including PC sensors and EID detectors that are different in size and therefore do not align perfectly, the gap between the PC sensors can be compensated for by the EID detector. That is, photons that fall into the gap and are not counted by the PC sensor array can be detected by the EID detector placed under the gap, and all photons of the incident X-rays can be detected by either the PC sensor or the EID detector. As a result, the image reconstructed from both PC and EID data has a higher resolution and a lower amount of noise compared to using only one of the PC and EID data. Furthermore, this arrangement reduces the reliance on the currently implemented tilted configuration. The tilted edges cause the shape of each detector element to deviate from a rectangle. With an asymmetric shape, the geometric center of the detector element is different from the signal weighting center, making it difficult to assign the proper geometric location of the detector element and align the corresponding parts of the adjacent detector modules. Furthermore, the pile-up problem is avoided with the EID detector, and since the sensor response is linear with respect to the incident X-ray flux, the EID data can be used to perform pile-up correction of the PC sensor data, resulting in a more accurate pile-up correction and a higher resolution.

[0006] An additional advantage of the proposed multi-layer CT detector is that it can reduce the cost of the silicon chips used for the PC sensor. Silicon has been demonstrated to have a good spectral response, making it a good candidate for the PC sensor. However, the depth of the silicon used for the PC sensor may have to be deeper than other semiconductor materials to ensure good dose efficiency and to ensure that the pile-up amount can be corrected. However, because some of the X-ray photons penetrating the CT detector can pass through the PC sensor and enter the EID detector, and the EID data can be used to more accurately correct the pile-up, the silicon depth can be reduced, reducing the cost and data size of the CT detector (fewer segments and channels). It also reduces the power consumption and heat dissipation of the CT detector.

[0007] Furthermore, if fast throughput is desired and conventional CT images are desired, as opposed to reconstructing images based on photon count values, the EID data alone can be used to generate conventional kVp images. Because the EID data projections are sufficiently filtered by the PC sensor layer, a narrower energy spectrum can be expected, which could result in improved performance in hardening radiation when images are reconstructed from the EID data projections alone.

[0008] These and other advantages and features of the present description will become readily apparent from the following detailed description, taken alone or in combination with the accompanying drawings. It is understood that the above summary is provided to introduce in a simplified form a selection of concepts that are further described in the detailed description. The above summary is not intended to identify key or essential features of the claimed subject matter, the scope of which is uniquely defined by the claims. Moreover, the claimed subject matter is not limited to implementations that solve the above disadvantages or disadvantages noted in other parts of this disclosure. [Brief description of the drawings]

[0009] The various aspects of the present disclosure can be further understood by reading the following detailed description and by reference to the drawings, in which: [Figure 1] FIG. 1 shows a diagram of an imaging system in accordance with one or more embodiments of the present disclosure. [Diagram 2] FIG. 1 is a block schematic diagram of an exemplary imaging system in accordance with one or more embodiments of the present disclosure. [Diagram 3] 1 is a schematic diagram of an exemplary detector array of a PCCT system in accordance with one or more embodiments of the present disclosure. [Figure 4] 1 illustrates an X-ray beam directed toward a multi-layer CT detector in accordance with one or more embodiments of the present disclosure. [Figure 5A] 1 is a schematic diagram of a multi-layer CT detector of a first aspect in accordance with one or more embodiments of the present disclosure. [Figure 5B] 1 is a schematic diagram of a multi-layer CT detector in a second view in accordance with one or more embodiments of the present disclosure. [Figure 6A] FIG. 2 is a schematic diagram of components of a two-dimensional PC sensor array of a multi-layer CT detector in accordance with one or more embodiments of the present disclosure. [Figure 6B] FIG. 1 is a schematic diagram illustrating the alignment of multiple two-dimensional PC sensor arrays of a multi-layer CT detector with multiple EID detectors of an underlying EID detector layer in accordance with one or more embodiments of the present disclosure. [Figure 6C] FIG. 6C is a schematic diagram showing another view of the PC sensor array and EID detector configuration of FIG. 6B in accordance with one or more embodiments of the present disclosure. [Figure 6D] FIG. 1 is a schematic diagram illustrating a first alternative configuration of a PC sensor array and an EID detector of a multi-layer CT detector in accordance with one or more embodiments of the present disclosure. [Figure 6E] FIG. 1 is a schematic diagram illustrating a second alternative configuration of a PC sensor array and an EID detector of a multi-layer CT detector in accordance with one or more embodiments of the present disclosure. [Figure 7] 1 is a flowchart illustrating an example method for enhancing the resolution of an image reconstructed from PC and EID data, in accordance with one or more embodiments of the present disclosure.

[0010] The drawings illustrate certain aspects of the described systems and methods. Together with the detailed description that follows, the drawings illustrate and explain the structures, methods, and principles described herein. In the drawings, dimensions of components may be exaggerated or distorted for clarity. Well-known structures, materials, or operations have not been shown or described in detail to avoid obscuring aspects of the described components, systems, and methods. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0011] DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS AND EMBODIMENTS OF THE SUBJECT MATTERS DISCLOSED HEREIN REGARD TO A METHOD AND SYSTEM FOR IMPROVING THE QUALITY OF IMAGES ACQUIRED BY A PHOTON COUNTING COMPUTED TOMOGRAPHY (PCCT) SYSTEM. Generally, in a computed tomography (CT) imaging system, an X-ray source emits a fan or cone beam toward an object (such as a patient). Generally, in a CT system, an X-ray source and a detector array rotate around the patient in an imaging plane in a gantry, and an image is generated from projection data of multiple views at different view angles. For example, one revolution of the X-ray source may generate 1000 views by the CT system. After being attenuated by the patient, the beam impinges on an array of radiation detectors. An X-ray detector or detector array typically includes a collimator for collimating the X-ray beam received by the detector, a scintillator disposed adjacent the collimator for converting the X-rays into light energy, and a photodiode for receiving the light energy from the adjacent scintillator and generating an electrical signal from the received light energy. The intensity of the attenuated beam radiation received by the detector array typically depends on the attenuation of the X-ray beam by the patient. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by it. The electrical signals are transmitted to and analyzed by a data processing system. The data processing system processes the electrical signals to generate an image.

[0012] Such conventional CT imaging systems utilize detectors that convert radiation energy into a current signal that is integrated over time, measured, and finally digitized. However, a drawback of such detectors is that they are unable to provide data or feedback regarding the number of photons detected and / or the energy of the photons. That is, the light emitted by the scintillator is a function of both the number of impinging x-rays and the energy level of the x-rays. Photodiodes cannot distinguish between the energy level and the number of photons from the scintillation. For example, two scintillators illuminated with equal intensity will result in equal output for each photodiode. However, despite producing equal light output, the number of x-rays received by each scintillator may differ, and further, the intensity of the x-rays may differ.

[0013] In contrast, PCCT detectors can perform photon counting and / or provide energy discrimination feedback to have high spatial resolution. PCCT detectors can operate in an X-ray counting mode, an energy measurement mode for each X-ray event, or both modes. Although multiple materials can be used to construct hybrid photon counting energy discriminating detectors, semiconductors have been shown to be one preferred material. Typical materials for such use include cadmium zinc telluride (CZT), cadmium telluride (CdTe), and silicon (Si), which can have cost-effective production capabilities. Other heavy semiconductors, such as thallium bromide (TlBr), mercury iodide (HgI), etc., can be used if they can be effectively produced in large quantities.

[0014] PCCT detectors support not only x-ray photon counting but also energy measurement or tagging, supporting the acquisition of both anatomical details and tissue characteristic information. In this regard, energy discrimination information or data can be used to mitigate effects such as radiation hardening. Furthermore, these detectors support the acquisition of tissue discrimination data, thus providing diagnostic information indicative of disease or other pathology. PCCT detectors can also be used to detect, measure, and characterize materials injected into the subject (such as contrast agents and / or other special materials) by using optimal energy weighting techniques that enhance the contrast of iodine and calcium (and other high atomic number materials). Contrast agents can include, for example, iodine injected into the bloodstream for improved visual visibility.

[0015] However, a drawback of direct conversion semiconductor detectors is that this type of detector cannot count at the x-ray photon flux typical of conventional CT systems. Also, saturation can occur at the detector when there is a thin section of the subject between the detector and the x-ray energy source or x-ray tube. The saturated region corresponds to the path of the thin section of the subject near or outside the width of the subject projected onto the detector fan arc. In many cases, the subject is roughly circular or elliptical in terms of the effect of attenuation of the x-ray flux and the effect of the attenuated incident intensity on the detector. In this case, the saturated region represents two separate regions at either end of the fan arc. In other, less typical but not rare, cases, saturation occurs in three or more separate regions elsewhere on the detector.

[0016] Pile-up is a phenomenon that occurs in PCT detectors when the source flux at the detector is so high that two or more x-ray photons deposit charge packets on one pixel close enough in time that the possibility of their signals interfering with each other cannot be ignored. There are two common types of pile-up phenomena, which have slightly different effects. In the first type, two or more events are separated in time so that they are recognized as separate events, but the signals overlap, reducing the accuracy of the energy measurement of the later arriving x-ray or both x-rays. This type of pile-up results in a reduction in the energy resolution of the system. In the second type of pile-up, two or more events arrive close enough in time that the system cannot separate them as separate events. In this case, the two or more events are recognized as one event with an energy that is the sum of the energies of the events, and the events are shifted towards higher energies in the spectrum. Pile-up also causes a significant reduction in counts at high x-ray fluxes, which in turn reduces the quantum efficiency (DQE) of the detector.

[0017] This pile-up can cause detector saturation at relatively low x-ray flux levels in direct conversion sensors. Beyond this level, the detector response becomes unpredictable, resulting in reduced dose utilization and loss of image information, resulting in noise and artifacts in x-ray projection and CT images. In particular, photon-counting direct conversion detectors saturate due to the intrinsic charge collection time (i.e., charge drift time) associated with each x-ray photon event. Pulse pile-up causes saturation when the x-ray photon absorption rate of each pixel is on the order of the inverse of this charge collection time.

[0018] PCT systems typically have one or more energy bins determined by a comparator that is part of the readout section of the data acquisition system (DAS). For one-bin systems, one energy threshold of the comparator is typically set at a high enough energy value so that there are few or no false noise counts, but low enough so that there is little loss of signal x-rays in the readout process. Such systems are subject to statistical errors and biases due to pile-up of multiple energy events, as will be explained.

[0019] Systems with many energy bins can be formed using multiple comparators in the readout DAS. Each comparator can be set to trigger on photons above a set level of energy, resulting in a register accumulating the number of photons above the corresponding x-ray energy level. The bin count values ​​can be weighted and summed to form a system output with a particular information content suitable for the imaging system. However, like one-bin systems, multi-bin systems are subject to degradation due to pile-up, resulting in a lower DQE. Although it is possible to correct for the average pile-up of the bin count values, statistical precision is lost. The signal-to-noise ratio (SNR) can be used to evaluate the weighted sum for the system output.

[0020] In some embodiments, the sensors of the PCCT detector can be configured as multiple sensor segments in a sensor array, with the sensor segments oriented in the direction of the incident x-rays. The sensor segments are used to prevent or account for pile-up behavior in the detector. The number of segments can be minimized to conserve application specific integrated circuit (ASIC) channels electrically coupled to the sensor array for analog-to-digital (A / D) conversion and readout. However, pile-up can occur within the sensor segments.

[0021] As disclosed herein, the inclusion of both PC and EID detectors provides the advantages of direct conversion detectors while minimizing pile-up motion and other disadvantages of PCCT systems. Specifically, disclosed herein is a multi-layer detector structure with an EID detector array disposed in a layer below the PC sensor array. The readout electronics for the PC sensor array and / or the EID detector array may be disposed below the EID detector array or to the side of the PC sensor array. In this disclosure, below refers to the location of the components relative to the incident x-rays, which are shown in Figures 4-6E as the x-rays descend vertically as they strike the components of the CT detector. Thus, the location of the EID detector array below or below the PC sensor array refers to the location of the EID detector array on the opposite side of the PC sensor array relative to the incident x-rays, and the location of the readout electronics below or below the EID detector array refers to the location of the readout electronics on the opposite side of the EID detector array relative to the incident x-rays (and PC sensor array).

[0022] Measurement of the count value of photons on the X-ray path is done using both the PC sensor array and the EID detector array, and the EID data and PC data are acquired simultaneously from the same X-ray path of the patient. To ensure that the EID detector array can obtain a sufficient signal, the PC sensor array can be configured to not capture all photons of the X-ray beam. In other words, the stopping power of the PC detector can be optimized by a thickness that is expected to reduce pile-up at the PC detector. For example, the stopping power can be 40% (e.g., 40% of the photons of the X-ray beam are detected by the PC sensor array, and 60% of the X-ray photons pass through the PC sensor array and are detected by the EID detector array). The X-ray photons that pass through the PC sensor array are captured by the EID detector array placed below the PC sensor array. Both the PC sensor and the EID detector can be used to generate spectral information as well as non-spectral information.

[0023] At least one advantage of the disclosed detector structure is that in addition to reducing pile-up motion in the PC detector, the EID data can be used to correct for pile-up motion in the PC data, as described in more detail below. In this manner, benefits from both types of detectors can be leveraged to improve image quality for a variety of clinical CT applications.

[0024] An example of a PCCT system that can be used to perform a contrast scan according to the present technology is shown in Figures 1 and 2. Figure 3 shows an example CT detector array of a PCCT system, where photons of X-rays directed by an X-ray source to a subject are counted by PCCT detectors of the PCCT detector array. The detectors can be multi-layer CT detectors including a PC sensor layer and an EID detector layer as shown in Figure 4. Two perspective views of an example multi-layer CT detector are shown in Figures 5A and 5B, where the readout electronics for the PC sensor layer and the EID detector layer are placed below the EID detector layer. In an alternative configuration of the multi-layer CT detector, a first readout electronics is used to count photons of the PC sensor layer and can be placed next to the PC sensor array as shown in Figure 6A. Multiple PC sensor arrays can be stacked on top of multiple EID detectors of the EID detector layer as shown in Figure 6B. A second readout electronics is used to measure the energy of the X-ray beam incident on the EID detector layer and can be placed below the EID detector layer as shown in Figure 6C. The PC sensor array can include multiple PC sensors on multiple silicon chips, and each PC sensor can be electronically coupled to a first readout electronics portion located beside the PC sensor array, as shown in FIG. 6D. In another alternative configuration, the first readout electronics portion can be located below the EID detector layer, and the multiple PC sensors are electronically coupled to the first readout electronics portion via one or more flexible cables routed around the EID detector layer, as shown in FIG. 6E. The resolution of an image reconstructed from both the PC and EID data can be improved by following one or more steps of the method described in FIG. 7.

[0025] 1-6E show exemplary configurations of relative positional relationships of various elements. When elements are shown as being in direct contact with or directly connected to one another, these elements can be referred to as directly contacting or directly connected elements, respectively, in at least one embodiment. Similarly, elements shown as being continuous or adjacent to one another can be referred to as being continuous or adjacent elements, respectively, in at least one embodiment. As an example, elements sharing a surface with one another can be referred to as being surface-sharing elements. As another example, elements positioned apart from one another can be referred to as being separate from one another, in at least one embodiment, if there is a space between the elements but no other elements are present. Also, as another example, elements shown above and below one another, opposite one another, or left and right from one another can be referred to as being above and below, opposite one another, or left and right elements, respectively. Furthermore, as shown in the figures, in at least one embodiment, the topmost element or the topmost point of an element can be referred to as the "top" of the element, and the bottommost element or the bottommost point of an element can be referred to as the "bottom" of the element. As used herein, top / bottom, upper / lower, above / below are relative to the vertical axis of the figure and are used to describe the location of elements of the figure relative to one another. Thus, elements shown on top of other elements, in one example, are vertically disposed on top of the other elements. As yet another example, the shapes of elements shown in the figures can be referred to as having those shapes (e.g., circular, rectilinear, flat, curvilinear, rounded, chamfered, angled, etc.). Furthermore, elements shown to intersect with one another, in at least one example, can be referred to as intersecting elements or elements intersecting one another. Furthermore, elements shown within or outside of another element can, in one example, be referred to as elements shown within or outside of another element.

[0026] FIG. 1 illustrates an exemplary PCCT system 100 configured to perform CT imaging using a photon-counting detector. In particular, the PCCT system 100 is configured to image a subject 112, such as a patient, an inanimate object, one or more manufactured parts, and / or a foreign object (such as an implant, a stent, and / or a contrast agent present in the body). In one embodiment, the PCCT system 100 includes a gantry 102, which may further include at least one X-ray source 104 that emits an X-ray radiation beam 106 (see FIG. 2) used to image the subject 112 lying on a table 114. In particular, the X-ray source 104 is configured to output the X-ray radiation beam 106 toward a detector array 108 disposed on an opposite side of the gantry 102 from the X-ray source 104. Although only a single X-ray source 104 is illustrated in FIG. 1, in an exemplary embodiment, multiple X-ray sources and multiple detectors may be used to output multiple X-ray radiation beams to obtain projection data at different energy levels corresponding to the patient. In some embodiments, the X-ray source 104 is capable of performing dual energy gemstone spectral imaging (GSI) with rapid kilovoltage peak (kVp) switching. In the embodiments described herein, the X-ray detector employed is a photon counting detector capable of distinguishing between X-ray photons of different energies.

[0027] In certain exemplary embodiments, the PCCT system 100 further includes an image processing unit 110 that reconstructs an image of the target volume of the subject 112 using an iterative image reconstruction method or an analytical image reconstruction method. For example, the image processor unit 110 can reconstruct an image of the target volume of the patient using an analytical image reconstruction method such as filtered back projection (FBP). As another example, the image processing unit 110 may reconstruct an image of the target volume of the subject 112 using an iterative image reconstruction method such as advanced statistical iterative reconstruction (ASIR), conjugate gradient (CG), maximum likelihood expectation maximization (MLEM), model-based iterative reconstruction (MBIR), etc. As described further herein, in some examples, the image processing unit 110 can also use analytical image reconstruction methods (such as FBP) in addition to the iterative image reconstruction methods.

[0028] In some CT imaging system configurations, an x-ray source emits a cone-shaped beam of x-ray radiation that is collimated in the XYZ plane of a Cartesian coordinate system and is commonly referred to as the "imaging plane." The x-ray radiation beam passes through the object being imaged (such as a patient or subject). After being attenuated by the object, the x-ray radiation beam impinges on an array of detector elements. The intensity of the attenuated x-ray radiation beam received at the detector array depends on the attenuation of the x-ray radiation beam by the object. Each detector element in the array produces a separate electrical signal that is a measurement of the x-ray beam attenuation at the detector location. Attenuation measurements from all detector elements are acquired separately to produce a transmission profile.

[0029] In some CT systems, the x-ray source and x-ray detector array are rotated by a gantry in an imaging plane around the object to be imaged so that the angle at which the x-ray beam intersects the object is constantly changing. A group of x-ray radiation attenuation measurements (e.g., projection data) obtained from the x-ray detector array at one gantry angle is called a "view." A "scan" of the object includes a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector.

[0030] 2 illustrates an exemplary imaging system 200 similar to the PCCT system 100 of FIG. 1. In accordance with aspects of the present disclosure, the imaging system 200 is configured to image a subject 204 (e.g., subject 112 of FIG. 1). In one embodiment, the imaging system 200 includes a detector array 108 (see FIG. 1). The detector array 108 further includes a plurality of detector elements 202 that sense an x-ray radiation beam 106 (see FIG. 2) passing through the subject 204 (e.g., a patient) and acquire corresponding projection data. In some embodiments, the detector array 108 is fabricated in a multi-slice configuration that includes multiple rows of cells or detector elements 202, and one or more additional rows of detector elements 202 can be arranged in parallel to acquire projection data.

[0031] In certain exemplary embodiments, imaging system 200 is configured to move to different angular positions about object 204 to acquire desired projection data. Thus, gantry 102 and the components mounted to gantry 102 may be configured to rotate about center of rotation 206, for example, to acquire projection data at different energy levels. Alternatively, in embodiments in which the projection angle relative to object 204 changes as a function of time, the mounted components may be configured to move along a general curve rather than along a circular arc.

[0032] As the X-ray source 104 and detector array 108 rotate, the detector array 108 collects data of the attenuated X-ray beam. The data collected by the detector array 108 is pre-processed and calibrated to adjust the data to represent the line integral of the attenuation coefficient of the scanned object 204. The processed data is commonly referred to as a projection. In some examples, the individual detectors or detector elements 202 of the detector array 108 may be photon counting detectors that record individual photon interactions into one or more energy bins.

[0033] The acquired set of projection data can be used for reference material discrimination (BMD). During BMD, the measured projections are transformed into a set of material density projections. The material density projections can be reconstructed to form a pair or set of material density maps or images of each reference material, such as bone, soft tissue, and / or contrast map. These density maps or density images can be correlated in turn to form a 3D volumetric image of the reference material (e.g., bone, soft tissue, and / or contrast agent) of the imaging volume.

[0034] Once reconstructed, the reference material image produced by the imaging system 200 reveals internal features of the subject 204 that are represented by the densities of the two reference materials. The density image may be displayed to show these features. In a conventional approach to diagnosing a medical event (such as a medical condition), or more generally, a radiologist or physician would review a hard copy or displayed density image to identify features of interest. Such features include lesions, the size and shape of particular anatomical structures or organs, and other features that may be discernible in the image based on the skill and knowledge of the individual practitioner.

[0035] In one embodiment, the imaging system 200 includes a control mechanism 208 that controls the movement of the components, such as the rotation of the gantry 102 and the operation of the x-ray source 104. In certain embodiments, the control mechanism 208 further includes an x-ray controller 210 configured to provide power and timing signals to the x-ray source 104. Additionally, the control mechanism 208 includes a gantry motor controller 212 configured to control the rotational speed and / or position of the gantry 102 based on imaging requirements.

[0036] In certain embodiments, the control mechanism 208 further includes a data acquisition system (DAS) 214 configured to sample analog data received from the detector elements 202 and convert the analog data into a digital signal for further processing. The DAS 214 may be further configured to selectively aggregate analog data from a subset of the detector elements 202 into a so-called macro-detector, as further described herein. The data sampled and digitized by the DAS 214 is transmitted to a computer or computing device 216. It is noted that in at least one embodiment, the computing device 216 may be the same as or similar to the image processing unit 110. In one embodiment, the computing device 216 stores the data in a storage device or mass storage device 218. The storage device 218 may be any type of non-transitory memory, such as a hard disk drive, a floppy disk drive, a compact disk read / write (CD-R / W) drive, a digital versatile disk (DVD) drive, a flash drive, and / or a solid-state storage drive.

[0037] Additionally, the computing device 216 provides instructions and parameters to one or more of the DAS 214, the X-ray controller 210, and the gantry motor controller 212 to control system operations (such as data acquisition and / or data processing). In certain exemplary embodiments, the computing device 216 controls system operations based on operator input. The computing device 216 accepts operator input including commands and / or scanning parameters, for example, by an operator console 220 operably coupled to the computing device 216. The operator console 220 may include a keyboard (not shown) or a touch screen to enable an operator to specify the commands and / or scanning parameters.

[0038] 2 illustrates one operator console 220, two or more operator consoles may be coupled to the imaging system 200, for example, to input or output system parameters, request exams, graph data, and / or view images. Additionally, in certain embodiments, the imaging system 200 may be coupled through one or more deployable wired and / or wireless networks (such as the Internet and / or virtual private networks, wireless telephone networks, wireless local area networks, wired local area networks, wireless wide area networks, wired wide area networks, etc.) to multiple displays, printers, workstations, and / or similar devices located locally or remotely within a facility or hospital or at entirely different locations.

[0039] In one embodiment, for example, imaging system 200 includes or is coupled to a Picture Archiving and Communication System (PACS) 224. In an exemplary embodiment, PACS 224 is further coupled to a remote system, such as a radiology department information system, a hospital information system, and / or is coupled to an internal or external network (not shown) to allow an operator at another location to provide commands and parameters and / or gain access to image data.

[0040] The computing device 216 may use operator-supplied and / or system-defined instructions and parameters to operate a table motor controller 226, which in turn may control the table 114, which may be a motorized table. In particular, the table motor controller 226 may move the table 114 to properly position the subject 204 in the gantry 102 to acquire projection data corresponding to a target volume of the subject 204.

[0041] As previously mentioned, the DAS 214 samples and digitizes the projection data acquired by the detector elements 202. The image reconstructor 230 then performs high speed reconstruction using the sampled and digitized x-ray data. Although FIG. 2 illustrates the image reconstructor 230 as a separate entity, in certain exemplary embodiments, the image reconstructor 230 may form part of the computing device 216. Alternatively, the image reconstructor 230 may not be present in the imaging system 200, and instead, the computing device 216 may perform one or more functions of the image reconstructor 230. Furthermore, the image reconstructor 230 may be located locally or remotely and may be operably connected to the imaging system 200 using a wired or wireless network. In particular, in one exemplary embodiment, computing resources in a "cloud" network cluster may be used for the image reconstructor 230.

[0042] In one embodiment, the image reconstructor 230 stores the reconstructed image in the storage device 218. Alternatively, the image reconstructor 230 may transmit the reconstructed image to the computing device 216 to generate patient information useful for diagnosis and evaluation. In some embodiments, the computing device 216 may transmit the reconstructed image and / or the patient information to a display or display device 232 communicatively coupled to the computing device 216 and / or the image reconstructor 230. In some exemplary embodiments, the reconstructed image may be transmitted from the computing device 216 or the image reconstructor 230 to the storage device 218 for short-term or long-term storage.

[0043] Referring now to FIG. 3, a CT detector array 300 is shown, which is a non-limiting example of the detector array 108 of FIG. 2. The detector array 300 includes rails 304 between which collimating blades or plates 306 are disposed. The plates 306 are disposed to collimate the x-rays 302, which, after being collimated, impinge on a number of detectors 308 of the detector array 300 disposed between the plates 306. As an example, the detector array 300 may include 57 detectors 308, each detector 308 having an array size of 64x16 pixel elements. As a result, the detector array 300 has 64 rows and 912 columns (16x57 detectors) and may simultaneously collect 64 slices of data per gantry rotation (e.g., the gantry 102 of FIG. 1).

[0044] As described in more detail below, each CT detector 308 may be a multi-layer CT detector that includes both a PC sensor array configured to directly convert radiation energy into an electrical signal that includes energy discrimination or photon count data, and an EID detector array configured to indirectly convert radiation energy into an electrical signal by means of a photodiode that receives light energy from an adjacently disposed scintillator.

[0045] 4 shows a proposed multi-layered CT detector structure 400, which includes a PC sensor array 402 arranged in a top layer and an EID detector array 404 arranged in a bottom layer (e.g., below the PC sensor array 402) for x-rays 406. The PC sensor array 402 is a two-dimensional sensor array including a first number M of PC sensors, and the EID detector array 404 is a two-dimensional sensor array including a second number N of EID detectors, where M can be greater than N. With M being greater than N, the PC sensors can have a first size (e.g., length and width) and the EID detectors can have a second size, where the second size is different from the first size. For example, a PC sensor 414 of the first number M of PC sensors can have a first length 422 and a first width 424. An EID detector 416 of the second number N of EID detectors can have a second length 432 and a second width 434 that are greater than the first length 422 and the first width 424.

[0046] The different sizes of each PC sensor and each EID detector can improve the quality of the reconstructed image. When M=N, the length and width of each sensor used for both M and N are the same. When the sizes of both the PC sensor and the EID detector are small, the EID detector may have a large dose efficiency problem. When the sizes of both the PC sensor and the EID detector are large, the PC sensor has a serious pile-up problem. When M>N, it is possible to simultaneously achieve good pile-up correction, high spatial resolution, and good dose efficiency.

[0047] The PC sensor 414 can have a first height 420 and the EID detector 416 can have a second height 430. In some embodiments, the first height 420 is greater than the second height 430. The first height 420 can be determined by a desired attenuation ratio or stopping power of incident photons at the PC sensor array 402. For example, when an X-ray beam 406 is incident on a PC sensor of the PC sensor array 402, a first portion of the photons of the X-ray beam 406 are detected by the PC sensor and a second portion of the photons of the X-ray beam 406 pass through the PC sensor and are detected by an EID detector of the EID detector array 404 located below the PC sensor. As the height of the PC sensor increases, a proportion of photons included in the first portion increases (e.g., more photons are detected by the PC sensor), and as the height of the PC sensor decreases, a proportion of photons detected by the first portion decreases. Thus, the height of the PC sensor (e.g., the first height 420) can be selected to achieve a desired attenuation ratio or stopping power of the PC sensor array 402.

[0048] The desired attenuation rate is based on the expected pile-up behavior in the PC sensor array 402. For example, as more photons are attenuated in the PC sensor array 402, the expected pile-up behavior in each PC sensor of the PC sensor array 402 increases. Thus, the first height 420 is selected to maintain the expected pile-up behavior below a desired threshold, but may not accurately correct for pile-up behavior above the threshold. Furthermore, the first height 420 may be selected based on the semiconductor material used in the PC sensor array 402, since semiconductor materials may have different attenuation rates. For example, when attempting to achieve a desired attenuation rate, if the semiconductor material is silicon, the first height 420 may be larger than if the semiconductor material is CZT or CdTe. The height of the EID detector may be selected to maximize the DQE to capture a majority (e.g., 95% or more) of the x-rays that pass through the PC sensor. For example, if the thickness of the PC detector is reduced to reduce pile-up, the EID detector may be thickened to capture a majority of the x-rays that pass through.

[0049] An additional benefit of having a PC sensor smaller than the EID detector is that if some of the X-ray beam enters the PC sensor array 402 through a boundary between each PC sensor 414, the photons of the X-ray beam will not be detected by any of the PC sensors 414. For example, a boundary 440 between the PC sensors of the PC sensor array 402 is located directly above a boundary 442 between the EID detectors of the EID detector array 404, and thus the X-ray beam can pass undetected through the PC sensor array 402 via the boundary 440 and through the EID detector array 404 via the boundary 442. However, in contrast to an M=N configuration in which the size of each PC sensor 414 is the same as the size of each EID detector 416 (e.g., in this configuration, the boundaries are aligned), in the M>N configuration shown in FIG. 4, photons of the X-ray beam entering the PC sensor array 402 through some of the boundaries between each PC sensor 414 can be detected by the EID detector 416 located below the boundary. For example, the X-ray beam passing through the boundary 444 between the PC sensors of the PC sensor array 402 can be detected by the underlying EID detector 416. As a result, the number of photons passing uncounted through both the PC sensor array 402 and the EID detector array 404 is reduced, resulting in a higher DQE and a higher quality reconstructed image. Furthermore, in some embodiments, the alignment of the PC sensor array 402 and the EID detector array 404 may be offset such that the boundary between each PC sensor 414 of the PC sensor array 402 is not located directly above the boundary between each EID detector of the EID detector array 404. In other words, by configuring the PC sensors to overlap, the number of X-ray beams passing through both the PC sensor array 402 and the EID detector 404 can be further reduced.

[0050] Additionally, if the PC sensor array 402 does not measure a signal due to the X-ray beam passing through a boundary between PC sensors 414 (or due to a bad pixel), the corresponding measurements from the EID detector array 404 can be used to help recover the missing data in the PC sensor array 402. Similarly, if the EID detector array 404 does not measure a signal due to the X-ray beam passing through a boundary between EID detectors 416, the corresponding measurements from a PC sensor 414 in the PC sensor array 402 that is positioned on that boundary can be used to help recover the missing data in the EID detector array 404.

[0051] 5A, a partial view of a multi-layer CT detector 500 of a PCCT system is shown. The multi-layer CT detector 500 may be a non-limiting embodiment of the detector 308 of FIG. Multiple rows of the multi-layer CT detector 500 may be arranged in a parallel configuration to form a detector array (e.g., detector array 108) for acquiring projection data, as described above.

[0052] 5A , the PC sensor array 502, the EID detector array 515, and the PCB 516 are arranged vertically relative to the direction of the incident x-ray beam 501 (indicated by arrow 507). The EID detector array 515 is disposed below the PC sensor array 502, and the PCB 516 is disposed below the EID detector array 515. In other embodiments, the PC sensor array 502, the EID detector array 515, and the PCB 516 may be arranged in different configurations relative to one another, as described in more detail below.

[0053] When rows of multi-layer CT detectors 500 are arranged in a parallel configuration to form a detector array, the rows of PC sensor array 502 and EID detector array 515 can form a two-dimensional sensor array similar to the PC sensor array 402 and EID detector array 404 described with reference to FIG. 4.

[0054] Both the PC sensor array 502 and the EID detector array 515 may be electronically coupled to an application specific integrated circuit (ASIC) 512 mounted on a PCB 516. As described in more detail below, the ASIC 512 may calculate the photon counts detected by the sensors of the PC sensor array 502 and digitize the integrated charge of the EID detector array 515. The PC sensor array 502 may be coupled to the ASIC 512 through a number of wire bonds 514 that run along a number of vias 513 and in the reflector material between the pixels of the EID detector array 515. The EID detector array 515 may be electronically coupled to the ASIC 512 through a number of wire bonds 519. Additionally, the PCB 516 may include connections 518 for the readout electronics of the detector 500, which may be used to calculate the number of photons detected by the PC sensor array 502 and digitize the charge signal of the EID detector array 515. The readout electronics may be external to the PCB 516 (eg, for use during assembly and testing). The readout electronics forms part of the DAS (eg, DAS 214) of the PCCT system.

[0055] The PC sensor array 502 can be configured to count photons that strike an edge 511 of the PC sensor array 502. In various embodiments, the PC sensor array 502 can be embedded in a chip made of a semiconductor material (such as silicon). The width of the chip can be one pixel at the edge 511 of the PC sensor array 502. Multiple sensors (such as sensor 414 in FIG. 4) can be embedded along the surface of the chip and extend along the length of the chip. Each sensor corresponds to one pixel 509 of the edge 511. Each sensor can include one or more sensor segments 505 embedded in columns 503 that extend below each pixel 509 and are oriented in the direction 507 of the incident x-ray beam 501 (e.g., vertical in FIG. 5A). Each sensor segment 505 in each column 503 can have a width of about one pixel across the surface of the PC sensor array 502, corresponding to a pixel 509 of the edge 511. Each sensor segment 505 in each column 503 is capable of counting the number of photons of the incident x-ray beam 501 that strike its edge 511 at a pixel 509 .

[0056] In other words, each column 503 may include multiple segments 505 stacked vertically in the column 503 in a direction 507. For example, each column 503 (shown vertically) may include a first segment at a first vertical position 504, a second segment at a second vertical position 506, and so on. In the embodiment shown in FIG. 5A, each column 503 includes two segments 505. In other embodiments, each column 503 may include two, three, four, or a different number of segments 505. The size of each segment 505 in a column 503 may be the same, or the size of each segment 505 in a column 503 may be different. For example, a first segment 505 in a column 503 may be smaller than a second segment 505 in the column 503 to reduce pile-up.

[0057] Each segment 505 may be electrically coupled to an ASIC 512 mounted on a PCB 516. In various embodiments, each segment 505 of the PC sensor array 502 may be electrically coupled to a sensor bond pad 510 of the PC sensor array 502 through a PCCT sensor trace 508. The sensor bond pad 510 may be electrically coupled to the ASIC 512 through a wire bond 514 that may extend through an EID detector array 515 and along a via 513 to the PCB 516.

[0058] Each segment 505 can detect the number of incident photons of the X-ray beam 501. When the X-ray beam 501 strikes the PC sensor array 502 at a pixel 509, the X-ray beam 501 can pass through the overlapped segments 505 of the corresponding column 503. As the X-ray beam 501 passes through each overlapped segment 505 of the column 503, the number of photons contained in the X-ray beam 501 can be detected at the associated segment 505.

[0059] For example, the exemplary X-ray beam 501 may be incident on a first segment 505 at a first vertical location 504 of the column 503, and the first segment 505 may detect a first number of photons of the exemplary X-ray beam 501. The first number of photons may be less than all of the photons of the exemplary X-ray beam 501, and a second number of the total photons may pass through the first segment 505 without being detected. The second number of photons of the exemplary X-ray beam 501 that pass through the first segment 505 (undetected) may be incident on a second segment 505 at a second vertical location 506 of the column 503. The second segment 505 may detect a third number of photons of the exemplary X-ray beam 501. The third number of photons may be less than the second number of photons, and a fourth number of photons may pass through the second segment 505 without being detected.

[0060] The number of photons detected in each vertically stacked segment 505 of column 503 may be different. A first portion of the photons of the exemplary X-ray beam 501 may be detected by the first segment 505, a second portion of the photons of the exemplary X-ray beam 501 may be detected by the second segment 505, and so on for additional segments. For example, many photons of the exemplary X-ray beam 501 may be detected in the first segment 505, fewer photons of the exemplary X-ray beam 501 may be detected in the second segment 505 than those detected in the first segment 505, and even fewer photons of the exemplary X-ray beam 501 may be detected in the third segment 505. Additionally, some photons of the exemplary X-ray beam 501 may not be detected in any of the segments 505 in column 503.

[0061] When the X-ray beam 501 passes through the PC sensor array 502, the X-ray beam 501 is attenuated to various degrees by the semiconductor material of the PC sensor array 502. When photons are detected in the first segment 505 and the second segment 505, the amount of energy of the X-ray beam 501 is reduced accordingly. When there are photons of the X-ray beam 501 that are not detected in the first segment 505 or the second segment 505 (and are not detected in the additional segments 505), the photons remaining in the X-ray beam 501 can be detected by the EID detector array 515 disposed below the PC sensor array 502. Specifically, the energy remaining in the X-ray beam 501 can be detected by the EID detector 530 disposed below the first segment 505 and the second segment 505 of the EID detector array 515.

[0062] Each EID detector 530 may include a scintillator 532 that converts x-rays into light energy, and a photodiode 534 disposed below the scintillator 532 that converts the light energy into an electrical signal. The electrical signal may be transmitted from a bond pad 521 of the photodiode 534 to the ASIC 512 via wire bonds 519. Each scintillator 532 of each EID detector 530 may be separated from one or more adjacent scintillators 532 of one or more adjacent EID detectors 530 by a cast reflector 536 to prevent light energy emitted in a corresponding scintillator 532 from mixing with one or more adjacent scintillators 532. While in FIG. 5A the EID detector 530 is shown as having the same width (e.g., one pixel) as each segment 505, it should be understood that in other embodiments the EID detector 530 may be wider than the sensors of the PC sensor array 502, as described with reference to FIG. 4.

[0063] By placing the EID detector 530 under the segment 505, all photons of the X-ray beam 501 can be detected by the multi-layer CT detector 500. A first portion of the photons of the X-ray beam 501 is detected by the PC sensor array 502, and a second portion of the photons of the X-ray beam 501 is detected by the scintillator 532. Because the EID detector 530 is placed under the first segment 505 and the second segment 505, the height of the PC sensor array 502 (e.g., and / or the height of the first segment 505 and the height of the second segment 505) can be configured such that a desired amount of X-ray flux is detected as photons by the PC sensor array 502, where the desired amount of X-ray flux is the number of photons that can be accurately corrected for pile-up motion. Thus, the height of the PC sensor array 502 can be lowered without affecting the quality of the resulting reconstructed image, which can reduce the cost of the multi-layer CT detector 500 since less semiconductor material is required. For example, if the semiconductor material is silicon, smaller and cheaper chips can be used.

[0064] When a photon strikes the segment 505, an analog electrical signal is generated and transmitted through the sensor trace 508 and the sensor bond pad 510 to the ASIC 512. The analog electrical signal is proportional to the amount of energy of the photon. The ASIC 512 can convert the analog electrical signal to a digital signal by counting the occurrence of photon hits with a counter. Furthermore, the ASIC can identify the energy imparted by the photon by comparing the amount of the electrical signal to one or more predetermined thresholds. Specifically, the ASIC 512 can include a plurality of comparators, each of which outputs a trigger signal to increment a corresponding digital counter by one when the analog signal exceeds a signal level threshold associated with the comparator. Each of the comparators of the plurality of comparators can have a different signal level threshold. For example, the ASIC 512 may include a first comparator with a first signal level threshold, a second comparator with a second signal level threshold higher than the first signal level threshold, a third comparator with a third signal level threshold higher than the second signal level threshold, and so on up to the maximum energy level of the spectrum of photons. The difference between the pairs of thresholds defines an energy range or energy bin. Thus, the number of photons whose energy falls within each bin can be recorded by the ASIC (or readout electronics). These photon count values ​​can be transmitted by the ASIC to the PCB over connections 524 for use in image reconstruction. Alternatively, the ASIC may first perform additional operations on the numeric count information (such as summing the individual photon count values ​​from the bins in a given column) to generate a total photon count value.

[0065] The advantage of including multiple segments in each column 503 is that pile-up behavior can be more accurately accounted for. X-rays are absorbed by the depth of the semiconductor material (e.g., silicon) regardless of how the semiconductor material is divided into segments. Each segment 505 acts as an independent sensing element that can generate a signal from the X-rays absorbed in the segment. As a result, each segment 505 counts the X-rays absorbed in the region of the semiconductor material associated with that segment 505, while other X-rays pass through the segment 505 unabsorbed. Each segment 505 is connected to its own counting channel circuit in the ASIC, so that the total X-ray count value is distributed to multiple channels. Each channel can then easily keep up with the absorption rate.

[0066] Although the effect of photon pile-up can be reduced by dividing each sensor of the PC sensor array 502 into segments, pile-up can still occur and cause inaccurate photon count values. However, by placing the EID detector array 515 under the PC sensor array 502, the size of the PC sensor array 502 and the size of the EID detector 515 can be advantageously selected to maximize the efficiency of pile-up correction and maximize the quality of the image reconstructed using the PC sensor data and the EID detector data. The advantages of the multi-layer CT detector 500 are that all electronic circuitry is under both detectors, limiting signal loss, and allowing smaller detector elements to be lined up to make a larger detector, which allows for a more compact design than wiring the elements of a larger detector elsewhere (e.g., next to the CT detector).

[0067] Also, placing the EID detector array 515 below the PC sensor array 502 advantageously allows data from the EID detector placed below the PC sensor to perform the pile-up correction applied to the photon count value of the PC sensor. Specifically, the incident flux detected by the EID detector placed below the PC sensor can be used to adjust the pile-up correction. The response of the EID detector is linear to all energies of the detected x-rays. Since the spectrum of the x-rays changes after passing through the patient and the PC detector, both the PC and EID can be used during pile-up calibration. When the x-ray flux is low, both the PC and EID data show a linear response to the incident x-rays. The PC data can be considered linear for all energy bins (energy spectrum). When the x-ray flux is high, the PC data may deviate from the linear response due to pile-up, but the EID data remains linear. Using both the energy spectrum of the PC data and the incident flux rate data of the EID, the pile-up calibration vector can be calculated using the pile-up model equation. Since the EID data is linear and independent of the x-ray flux, the EID data improves the quality of the pile-up calibration. A pile-up correction is performed by applying the pile-up calibration vector to the patient data.

[0068] FIG. 5B illustrates a side perspective view 550 of the multi-layer CT detector 500 of FIG. 5A. The side perspective view 550 illustrates a plurality of PC sensor arrays disposed on a two-dimensional EID detector array 551. Each PC sensor array of the plurality of PC sensor arrays can be the PC sensor array 502 of the non-limiting example of FIG. 5A, and the EID detector array 551 can be the EID detector array 515 of the non-limiting example of FIG. 5A. The plurality of PC sensor arrays includes a first PC sensor array 552, a second PC sensor array 554, a third PC sensor array 556, and a fourth PC sensor array 558. The EID detector array 515 includes a plurality of EID detectors (such as a first EID detector 560 and a second EID detector 562). The first EID detector 560 can be separated from the second EID detector 562 by a cast reflector 564 (e.g., the cast reflector 536 of FIG. 5A). It should be understood that side perspective 550 may include a portion of multi-layer CT detector 500 that may include additional PC sensor arrays and / or EID detectors without departing from the scope of this disclosure.

[0069] FIG. 5B illustrates the relative thicknesses of different components of the multi-layer CT detector 500. In FIG. 5B, each EID detector of the plurality of EID detectors can have a different length, width, and / or thickness than each PC sensor array of the plurality of PC sensor arrays. For example, a thickness 570 of the first PC sensor array 552, the second PC sensor array 554, the third PC sensor array 556, and the fourth PC sensor array 558 can be thinner than a thickness 572 of the first EID detector 560 and the second EID detector 562. Because the thickness 570 is thinner than the thickness 572, the plurality of PC sensor arrays can be aligned with various plurality of EID detectors of the EID detector array 551. For example, the first PC sensor array 552 and the second PC sensor array 554 can be disposed above the first EID detector 560, and the third PC sensor array 556 and the fourth PC sensor array 558 can be disposed above the second EID detector 562. As a result, an incident X-ray beam 580 incident on a PC sensor of the first PC sensor array 552 is partially attenuated in the first PC sensor array 552, and a weakened X-ray beam 580 (e.g., an X-ray beam with reduced energy) can pass through the first PC sensor array 552 and be detected by the first EID detector 560. Similarly, an incident X-ray beam 582 incident on a PC sensor of the second PC sensor array 554 is partially attenuated in the second PC sensor array 554, and a weakened X-ray beam 582 can pass through the second PC sensor array 554 and be detected by the first EID detector 560. In contrast, an incident X-ray beam 584 incident on the PC sensor of the third PC sensor array 556 is partially attenuated within the third PC sensor array 556, and the weakened X-ray beam 584 passes through the third PC sensor array 556 and can be detected by the second EID detector 562, and an incident X-ray beam 586 incident on the PC sensor of the fourth PC sensor array 558 is partially attenuated within the fourth PC sensor array 558, and the weakened X-ray beam 586 passes through the fourth PC sensor array 558 and can be detected by the second EID detector 562.

[0070] In the illustrated embodiment, each of the first PC sensor array 552, second PC sensor array 554, third PC sensor array 556, and fourth PC sensor array 558 is electrically coupled to the ASIC 512 through a corresponding wirebond 522 (e.g., through a wirebond through a path or via in a cast reflector 564). Each of the first EID detector 560 and second EID detector 562 is also electrically coupled to the ASIC 512 through a corresponding wirebond 519. In other embodiments, one or more additional ASICs 512 may be mounted on the PCB 516, where portions of the PC sensor arrays and / or portions of the EID detector arrays may be electronically coupled to the one or more additional ASICs 512.

[0071] One advantage of this configuration is that X-rays traveling between the PC sensors or between the PC sensor arrays can be detected by the EID detectors located below the PC sensor arrays. For example, an X-ray beam 590 may enter the multi-layer CT detector 500 between the first PC sensor array 552 and the second sensor array 554, and the X-ray beam 590 may not be detected by either the first PC sensor array 552 or the second sensor array 554. However, the X-ray beam 590 may enter and be detected by the EID detector 560 located below the first PC sensor array 552 and the second sensor array 554. In this manner, since the EID detectors 560 and 562 overlap the PC sensor arrays 552, 554, 556, and 558, the X-ray beams passing undetected through the PC sensor layers of the multi-layer CT detector 500 can be detected by the EID detector arrays of the multi-layer CT detector 500 to obtain a high-quality reconstructed image.

[0072] 5A and 5B is that by locating the electronics (e.g., PCB 516, ASIC 512, etc.) below the EID detector array, the length of the circuitry and electronic connections between the PC sensor array, the EID detector, and other elements are reduced, resulting in a compact CT detector with less signal loss and reduced material costs than other configurations. Locating the electronics below the EID detector array advantageously allows smaller detector elements to be collocated to create a larger detector in a cost-effective manner.

[0073] 5A and 5B illustrate various advantages of a CT detector including an EID detector array disposed below a PC sensor array, where the electronics, including PCB 516, ASIC 512, and other readout electronics, are disposed in a vertical arrangement below the EID detector array and the PC sensor array. In contrast, FIGS. 6A, 6B, 6C, 6D, and 6E illustrate alternative configurations of a multi-layer CT detector including an EID detector array disposed below a PC sensor array, where some of the electronics are not disposed below the EID detector array.

[0074] FIG. 6A shows an alternative configuration 600 of a PC sensor array in a CT detector, in which electronic circuitry including an ASIC used to count photons of an incident X-ray beam and measure the energy of the photons is located on one side of the sensor array. In the alternative configuration 600, a PC sensor array 606 is formed by attaching a silicon chip 602 to a flexible or ceramic PCB 604. An ASIC 616 is attached to the PCB 604 on one side of the silicon chip 602. The silicon chip 602 can be divided into multiple PC sensors 610, each of which can include one or more sensor segments, as described above with reference to FIG. 5A. Each PC sensor 610 can include solder points 612 by which the PC sensor 610 can be coupled to a corresponding circuit 614 embedded in the PCB 604 that electronically connects the PC sensor 610 to the ASIC 616. A perspective view 608 shows the top of the PC sensor array 606, with the silicon chip 602 on the front side 622 of the PC sensor array 606 attached to a PCB 604 on the back side 620 of the PC sensor array 606. In contrast to the CT detector shown in Figures 5A and 5B, the illustrated ASIC 616 is located to one side of the PC sensor array 606, rather than below the PC sensor array 606 (e.g., below the EID detector array).

[0075] 6B shows a first perspective view 650 of a CT detector including a plurality of PC sensor arrays, each of which may be the non-limiting example PC sensor array 606 of FIG 6A. The plurality of PC sensor arrays includes a first PC sensor array 651, a second PC sensor array 652, a third PC sensor array 653, a fourth PC sensor array 654, a fifth PC sensor array 655, and a sixth PC sensor array 656, which are arranged in a parallel configuration to create a two-dimensional array of individual PC sensors, as described above with reference to FIG 4. A first perspective view 650 illustrates the alignment of multiple PC sensor arrays with multiple EID detectors of an EID detector array 659, where a first PC sensor array 651, a second PC sensor array 652, and a third PC sensor array 653 can be positioned above the EID detectors of a first EID detector row 657, and a fourth PC sensor 654, a fifth PC sensor 655, and a sixth PC sensor 656 can be positioned above the EID detectors of a second EID detector row 658. Thus, a first portion 660 of a PC sensor (e.g., PC sensor 610) included in the first PC sensor array 651 can be positioned above a first EID detector 663 of the first EID detector row 657, a second portion 661 of the PC sensor included in the first PC sensor array 651 can be positioned above a second EID detector 664 of the first EID detector row 657, and a third portion 662 of the PC sensor included in the first PC sensor array 651 can be positioned above a third EID detector 665 of the first EID detector row 657. Similarly, a first portion 660 of the PC sensor included in the second PC sensor array 652 and the third PC sensor array 653 can be positioned above the first EID detector 663, a second portion 661 of the PC sensor included in the second PC sensor array 652 and the third PC sensor array 653 can be positioned above the second EID detector 664, and a third portion 662 of the PC sensor included in the second PC sensor array 652 and the third PC sensor array 653 can be positioned above the third EID detector 665.

[0076] Figure 6C shows a second perspective 670 of the CT detector of Figure 6B, which shows a side view of an EID detector aligned below the PC sensor array with respect to the downward x-ray beam 601. Figure 6C shows a PC sensor array 606 (such as PC sensor arrays 651, 652, 653, 654, 655, and 656) that is disposed above a first EID detector 672, a second EID detector 673, and a third EID detector 674 of an EID detector array 677. For example, the first EID detector 672 corresponds to the EID detector 663 of Figure 6B, the second EID detector 673 corresponds to the EID detector 664, and the third EID detector 674 corresponds to the EID detector 665. The photodiode layer 675 of the CT detector can include multiple photodiodes (e.g., photodiode 534 of FIG. 5A) that convert photons detected by the first EID detector 672, the second EID detector 673, and the third EID detector 674 into electrical signals that can be sent to the second ASIC 676 of the CT detector.

[0077] For example, the X-ray beam 601 is incident on the PC sensor 671 of the PC sensor array 606. A portion of the total number of photons of the X-ray beam 601 can be detected by the PC sensor 671. The detected portion of photons can be counted by the ASIC 616. If the X-ray beam 601 includes photons that are not detected, the X-ray beam 601 can pass through the PC sensor 671 and be incident on the EID detector 672. However, as a result of some of the photons being detected by the PC sensor 671, a reduced portion of the X-ray beam flux 601 is incident on the EID detector 672. The EID detector 672 can convert the energy of the X-ray beam 601 into light, which can cause the photodiodes of the photodiode layer 675 to generate a current. The current can be sent to a second ASIC 676 of the CT detector, which can be located below the photodiode layer 675. Both the flux-reduced X-ray beam 601 detected by the photodiode and the photon count values ​​of the portion of the photons detected by the PC sensor 671 can be used to reconstruct an image.

[0078] By reconstructing an image based on data from both the PC sensor 671 and the EID detector 672, the resolution of the image can be as high as an image reconstructed by a PCCT detector (e.g., higher than the resolution achievable by an EID detector alone) and the noise level of the image can be lower than the noise level of an image reconstructed by a PCCT detector. The resolution is higher because, among other factors described herein, the PC sensor 671 is smaller than the PC sensor of a PCCT detector without an EID detector layer, and data from the EID detector 672 is used to assist in the pile-up correction applied to the photon count values ​​from the PC sensor 671.

[0079] When estimating spectral information, the PC data of multiple energy bins from the PC sensor 671 can be used. However, the EID data from the EID detector 672 can also be used. The EID data can be treated as wide-bin independent responses. Through system response modeling and / or phantom calibration, the EID data as well as the PC data can be used as part of the material decomposition process.

[0080] Additionally, the EID data itself can be used to generate conventional kVp images without the use of PC data. This can be clinically advantageous when fast throughput is desired and conventional CT images are desired. Because the EID projections are filtered by the PC sensor, a narrow energy spectrum can be expected, resulting in improved linear curing performance when images are reconstructed from the EID projections alone. To reduce noise in images generated from the EID data, photon count values ​​from the PC sensor can be combined with the EID data to generate conventional kVp images. Combining the photon count values ​​from the PC sensor with the EID data can include assigning energy weightings to the PC sensor data and the EID data.

[0081] FIG. 6D illustrates an alternative configuration 680 of the CT detector of FIG. 6C, where the PC sensor array 681 includes multiple silicon chips 602 mounted on a single PCB 688 that includes an ASIC 616 disposed beside the PC sensor array 681. For example, the PC sensor array 681 includes a first silicon chip 682 including multiple sensors 610, a second silicon chip 684 including multiple sensors 610, and a third silicon chip 686 including multiple sensors 610, where each sensor 610 of the first silicon chip 682, the second silicon chip 684, and the third silicon chip 686 is electronically coupled to the ASIC 616. In some embodiments, one or more additional ASICs 616 may be included on the PCB 688. Each sensor 610 of the first silicon chip 682, the second silicon chip 684, and the third silicon chip 686 may be disposed above the EID detectors of the EID detector array 689, as described above.

[0082] FIG. 6E illustrates a CT detector 690 with a similar configuration to that shown in FIG. 6C, but with a different semiconductor material (e.g., a material other than silicon) used for the PC sensor array 691. For example, the different semiconductor material may be Cadmium Zinc Telluride (CZT), Cadmium Telluride (CdTe), or a different type of semiconductor material. Similar to the PC sensor array 606 of FIG. 6A, the PC sensor array 691 includes a plurality of PC sensors (e.g., sensor segments) 694, each of which is electronically coupled to an ASIC 696 configured to count photons detected at the plurality of PC sensors 694. The density of the different semiconductor material may be greater than that of silicon, and each of the PC sensors 694 may have a height (relative to the downward X-ray beam 601) that is less (e.g., 1-3 mm) than the height (e.g., 10 mm-50 mm) of the silicon chip 602.

[0083] 6A-6D, the ASIC 696 is disposed below the EID detector array 697 (e.g., EID detector array 677), which includes a plurality of EID detectors 698 disposed above the photodiode layer 699, as described above. The PC sensor array 691 can be electronically coupled to the ASIC 696 through one or more flexible cables 693, which allows the ASIC 696 to be disposed below the photodiode layer 699. By disposing the ASIC 696 below the EID detector array 697, the amount of material between the PC sensor array 691 and the EID detector array 697 can be minimized. Because the material between the PC sensor array 691 and the EID detector array 697 absorbs x-rays, minimizing the amount of material between the PC sensor array 691 and the EID detector array 697 increases the efficiency of the CT detector 690, resulting in a higher quality reconstructed image. A second ASIC 695 may be disposed below the photodiode layer 699 and may be used to measure the amount of detected energy of the x-ray beam passing through the PC sensor array 691 and impinging on an EID detector 698 of the EID detector array 697. Thus, each image reconstructed based on the x-rays detected by the CT detector 690 may be reconstructed using the photon count values ​​detected by the PC sensor array 691 and counted by the ASIC 696, and the x-ray beam energy measured by the EID detector array 697 and integrated by the ASIC 695.

[0084] 7, a flow chart is shown. The flow chart illustrates an exemplary method 700 for increasing the resolution of an image reconstructed from X-rays incident on a multi-layer CT detector (e.g., multi-layer CT detector 500) of a CT system using PC data received from a PC sensor of a PC sensor array (e.g., PC sensor array 402) of the multi-layer CT detector and EID data received from an EID detector of an EID detector array (e.g., EID detector array 404) of the multi-layer CT detector. As described above, by reconstructing an image based on both PC data and EID data, more X-ray photons can be detected and the EID data can be used to assist in pile-up correction of the PC data, resulting in an image with higher resolution and less noise than an image reconstructed from only the PC sensor or the EID detector. Additionally, as described below, the method 700 can be used to further increase the resolution using a trained convolutional neural network. The method 700 can be stored as instructions in a non-transitory memory and executed by one or more processors of a computing device of a CT imaging system (e.g., computing device 216 of imaging system 200 of FIG. 2).

[0085] Method 700 begins at 702. Method 700 includes beginning a scan of a subject using a multi-layer CT detector. The multi-layer CT detector may include a first layer of PC sensors overlaid on a second layer of EID detectors, as described above with reference to FIG. 4. In some embodiments, the multi-layer CT detector may be a non-limiting example of the multi-layer CT detector 500 of FIGS. 5A and 5B or the multi-layer CT detector 690 of FIG. 6E, where the readout electronics of the multi-layer CT detector are disposed below the second layer of EID detectors relative to the direction of the incident x-rays. In other embodiments, the multi-layer CT detector may be a non-limiting example of the multi-layer CT detector described with reference to FIGS. 6A-6D, where a first readout electronics configured to count photons at each PC sensor array of the first layer of PC sensors is disposed to one side of the PC sensor arrays, and a second readout electronics configured to measure the energy of the x-ray beam incident on the EID detectors of the second layer of EID detectors is disposed below the second layer of EID detectors.

[0086] At 704, the method 700 generates an image I based on signals received from a first layer of PC sensors of the multi-layer CT detector after pile-up correction of the PC data using both the PC data and the EID data. 1 The signal may be transmitted from a plurality of PC sensor arrays (such as PC sensor array 606 in FIG. 6A) in a first layer of PC sensors to corresponding readout electronics, such as an ASIC (e.g., ASIC 512 or ASIC 616).

[0087] At 706, the method 700 generates an image I based on a signal simultaneously received from a second layer of the EID detector of the multi-layer CT detector. 2The image I may include a second set of images (e.g., the images are generated from the same set of x-ray beams). Signals may be transmitted from the multiple EID detectors (e.g., EID detectors 663, 664, and 665 in FIG. 6B) in the second layer of EID detectors to readout electronics (e.g., ASIC 512 or ASIC 676 in FIG. 6C) corresponding to the EID detectors. Since the first size of the PC sensor is smaller than the second size of the EID detectors, the image I may include a second set of images (e.g., the images are generated from the same set of x-ray beams). 2 Here is the image I 1 It can have a lower resolution than 1 and I 2 Since both of them use a part of the incident X-ray signal, I 1 and I 2 may be noisy.

[0088] At 708, the method 700 generates a low-resolution image I 2 High resolution image I 1 The process involves training a deep learning CNN to map the input image pairs (I 2 ,I 1 ) is generated. Image I 2 is the input image, and image I 1 is the ground truth image. During CNN training, each input image I 2 For each image, the weights of the various layers of the CNN are adjusted by a gradient descent process using backpropagation, and the CNN output is compared with the corresponding ground truth image I 1 After training the CNN, when a new reconstructed image is input to the trained CNN, the trained CNN can output a high-resolution image with the increased resolution of the new reconstructed image.

[0089] At 710, the method 700 performs a step of: 2 ,I 1 ) to fused image I f In various embodiments, the fusion image I f Here is the image I 1 and I 2where I f =aI 1 +bI 2 It is. f To minimize the amount of noise in f The parameters a and b can be adjusted to maximize the resolution of I, or to achieve a target resolution and a target noise level. Once this fusion step is performed, the fused image I f I 1 and I 2 It has a lower amount of noise than both I 2 Although it has a higher resolution than I 1 The target noise level can be achieved by using commonly used methods. For example, a water phantom is scanned and the image I 1 and I 2 Since the image of the water phantom is uniform, I 1 and I 2 Noise level and fusion image I f =aI 1 +bI 2 In order to minimize or achieve a certain noise level in the output image, appropriate values ​​of a and b are determined, and the output fusion image I f =aI 1 +bI 2 The noise of image I is the minimum noise or a certain level of noise. 1 Resolution and Image I 2 Scan and image a resolution phantom using tungsten wires to obtain a specific resolution between the resolution of I 1 and Image I 2 and for given a and b, we reconstruct the fused image I f By finding the appropriate values ​​of a and b, the resolution of I f A target resolution of 1000x ... can be achieved.

[0090] At 712, the method 700 generates a fusion image I f The denoised fused image I is obtained by denoising fdThe fusion image I f is the denoised image I fd In the weighted average denoising method, the output value of pixel (i,j) is calculated by the following formula: where w is the weight of the neighboring pixels and N is used to define the neighborhood size. Usually, w is the weight of I f The weights can be calculated using various methods (e.g., bilateral filtering or non-local means).

number

[0091] At 714, the method 700 generates a denoised fusion image I fdis fed into the trained deep learning CNN to generate the denoised fused image I fd The inference output image I has a higher resolution than output This includes generating fd Since the amount of noise has been reduced as a result of step 712, the noise does not increase too much in the trained CNN during the inference phase. fd By feeding this to a trained deep learning CNN, we obtain the output image I output can have a high resolution comparable to that which would be obtained with a photon counting CT system, but with low noise.

[0092] At 716, the method 700 includes displaying the output image on a display device of the CT system. The method 700 then ends.

[0093] In another embodiment, the image pair (I 2 ,I 1 ) each image I 2 and I 1 can be denoised (e.g., by traditional weighted averaging, guided noise reduction, or deep learning noise reduction in step 712) before the deep learning CNN is trained. The CNN then uses the denoised image pair (I 2d ,I 1d ) can be trained on a set of images I 2d is the input image, and image I 1d is the ground truth image. The denoised image pair (I 2d ,I 1d ) and the result of training the CNN on the original image pair (I 2 ,I 1), during a subsequent inference phase (e.g., with new reconstructed images based on PC and EID data), the trained CNN can output images with higher resolution and less increased noise. The images output by the trained CNN can then be fused as described in step 710. The parameters a and b can be adjusted such that the resulting fused image has a further increased resolution and a further reduced noise.

[0094] Thus, the multi-layer CT detector configuration including a first PC sensor layer and a second EID detector layer provides a new approach to reconstruct images from CT scan data that leverages the advantages of both PC and EID detectors to achieve high image quality for various clinical CT applications including spectral and non-spectral tasks. The combination of the two techniques overcomes various challenges to generate high quality reconstructed images. By placing the second EID detector layer below the first PC sensor layer and the readout electronics below the second EID detector layer, data size can be reduced and thermal and power management can be improved. By including the second EID detector layer below the first PC sensor layer, the height of the PC sensors in the PC sensor layer can be reduced, reducing the size and cost of the silicon chips used in the PC sensor layer. In some embodiments, the cost of the multi-layer CT detector can be further reduced by placing the first readout electronics (e.g., electronics for photon counting sensors) next to the PC sensor array and the second readout electronics (e.g., electronics for EID data) below the EID detector layer.

[0095] Because the second EID detector layer can detect the energy level of the incident X-rays that have passed through the PC sensor layer, the PC sensor layer can include a smaller size PC sensor than the EID detector, resulting in a higher resolution of the resulting image. The EID data can also be used to increase the accuracy of the pile-up correction applied to the PC sensor data, resulting in higher resolution and lower noise. Another advantage of having an EID detector that is larger than the PC sensor is that X-rays that pass through the gaps between the boundaries of adjacent PC sensors can be detected by the underlying EID detector, thereby increasing the number of X-rays and the overall amount of photons detected by the multi-layer CT detector. If high speed processing is desired, the EID detector data can be used to reconstruct an image without using the PC sensor data. Here, the EID detector data can be more accurate than a single-layer EID detector due to the filtering effect of the PC sensor layer, which can harden the X-ray beam.

[0096] As a result of the various advantages of the multi-layer CT detector, it is possible to reconstruct images with higher resolution than the EID detector, comparable resolution to the PCCT detector, and less noise than the PCCT detector. Furthermore, a method is provided to further improve the resolution of the images generated by the PC sensor data and the EID detector data without increasing the amount of noise using a CNN trained with an image pair containing an input image from the EID detector data and a target image from the PC sensor data. By having both the PC sensor and the EID sensor on one system and measuring the same object simultaneously, corrections can be made in the image chain, producing images with high spatial resolution and low noise, supporting the usual clinical needs of CT systems while providing many advanced clinical applications.

[0097] The technical effect of a multi-layer CT detector including a first layer of PC sensors and a second layer of EID detectors positioned below the PC sensors is that it has a higher resolution than an EID detector, a comparable resolution to a PCCT detector, and can reconstruct images with less noise than a PCCT detector.

[0098] The present disclosure provides support for a computed tomography (CT) detector system. The CT detector system includes a layer of energy integrating detectors (EIDs) arranged on an opposite side to an incident x-ray from a layer of photon counting (PC) sensors, the number of the PC sensors being greater than the number of the EID detectors. In a first embodiment of the system, the system further includes an image processing unit configured to correct the PC data using the EID data, denoise an image reconstructed from the EID data and the PC data using a deep learning convolutional neural network (CNN) trained with a plurality of pairs of images, and increase the resolution of the reconstructed image, each pair of images including a target image reconstructed from a first signal from the layer of PC sensors and an input image reconstructed from a second signal from the layer of EID detectors, the EID data and the PC data being acquired simultaneously from the same patient line path. In a second embodiment of the system, the second embodiment optionally includes the first embodiment, and both the PC data and the EID data are used to generate spectral and non-spectral information. In a third embodiment of the system, the third embodiment optionally includes the first or second embodiment, and the image is reconstructed from EID data rather than from PC data. In a fourth embodiment of the system, the fourth embodiment optionally includes one or more of the first to third embodiments, and boundaries between each PC sensor of the plurality of PC sensors are not aligned with boundaries between each EID detector of the plurality of EID detectors. In a fifth embodiment of the system, the fifth embodiment optionally includes one or more of the first to fourth embodiments, and further includes readout electronics disposed on an opposite side of the layer of EID detectors with respect to incident x-rays, and signals from each PC sensor of the plurality of PC sensors are transmitted to the readout electronics through wire bonds routed between the EID detectors.In a sixth embodiment of the system, the sixth embodiment optionally includes one or more of the first to fifth embodiments, further comprising a first set of readout electronics and a second set of readout electronics, wherein a signal from each EID detector of the plurality of EID detectors is transmitted to the first set of readout electronics and a signal from each PC sensor of the plurality of PC sensors is transmitted to the second set of readout electronics. In a seventh embodiment of the system, the seventh embodiment optionally includes one or more of the first to sixth embodiments, wherein the first set of readout electronics and the second set of readout electronics are disposed on opposite sides of the layer of the EID detector with respect to the incident x-rays, and a signal from each PC sensor of the plurality of PC sensors is transmitted to the second set of readout electronics through one or more flexible cables routed around the layer of the EID detector. In an eighth embodiment of the system, the eighth embodiment optionally includes one or more of the first to seventh embodiments, wherein a first set of readout electronics is disposed on an opposite side of the EID detector layer relative to the incident x-rays, and a second set of readout electronics is disposed on one side of the layer of PC sensors. In a ninth embodiment of the system, the ninth embodiment optionally includes one or more of the first to eighth embodiments, wherein the PC sensors of the layer of PC sensors are arranged in the form of a two-dimensional PC sensor array, each PC sensor array including a semiconductor material mounted on a printed circuit board including a readout electronics unit of the second set of readout electronics and a plurality of circuits, each circuit of the plurality of circuits electrically coupling a PC sensor of the PC sensor array to the readout electronics. In a tenth embodiment of the present system, the tenth embodiment optionally includes one or more or each of the first to ninth embodiments, and the image processing unit is configured to correct the PC data using the EID data by adjusting a pile-up correction applied to the photon count values ​​detected by the PC sensors of the layer of PC sensors based on both the energy spectrum of the PC data and the incident flux rate detected by an EID detector positioned on the opposite side of the layer of PC sensors relative to the incident X-rays.In an eleventh embodiment of the present system, the eleventh embodiment optionally includes one or more or each of the first to tenth embodiments, and further includes: increasing the resolution of an image reconstructed from EID data and PC data using a deep learning CNN trained on multiple pairs of images including a target image reconstructed from a first signal from the PC sensor layer and an input image reconstructed from a second signal from the EID detector layer; creating a fusion image from the target image and the input image; and removing noise from the fusion image using at least one of a weighted average method, a guided noise reduction method, and a deep learning noise reduction method. The method further includes inputting the denoised fusion image to the trained CNN to generate an output image, the output image having a higher resolution than the denoised fusion image.

[0099] The present disclosure also provides support for a method for a computed tomography (CT) system. The method includes reconstructing a first image based on photon counting (PC) data received from a PC sensor layer of a CT detector of the CT system, reconstructing a second image based on energy integral (EI) data received from an EID detector array of the CT detector, training a convolutional neural network (CNN) using the first image as ground truth data such that the second image has a higher resolution, combining the first image and the second image to generate a fusion image, inputting the fusion image into the trained CNN to generate an output image, the output image having a higher resolution and lower noise than the fusion image, and displaying the output image on a display device of the CT system. In a first embodiment of the method, the EID detector array is disposed on an opposite side of the PC sensor layer with respect to incident x-rays. In a second embodiment of the method, the second embodiment optionally includes the first embodiment, and a boundary between PC sensors of the PC sensor layer is not aligned with a boundary between EID detectors of the EID detector array. In a third embodiment of the method, the third embodiment optionally includes the first or second embodiment, and combining the first image and the second image to generate the fusion image further includes weighting the first image with a first parameter and weighting the second image with a second parameter. In a fourth embodiment of the method, the fourth embodiment optionally includes one or more of the first to third embodiments, and the first parameter and the second parameter are selected to one of minimize an amount of noise in the fusion image and maximize a resolution of the fusion image. In a fifth embodiment of the method, the fifth embodiment optionally includes one or more of the first to fourth embodiments, and the method further includes denoising the fusion image using at least one of a weighted average method, a guided noise reduction method, and a deep learning noise reduction method.In a sixth embodiment of the method, the sixth embodiment optionally includes one or more or each of the first to fifth embodiments, and the method further includes performing pile-up correction on photon count values ​​of the PC data received from the PC sensor layer, the pile-up correction being performed based on the EID data received from the EID detector array.

[0100] The present disclosure also provides support for a computed tomography (CT) detector including a first number of photon counting (PC) sensors disposed in a first layer of the CT detector, a second number of energy integrating detectors (EIDs) disposed in a second layer of the detector, the second layer being on an opposite side of the first layer with respect to incident x-rays, the second number being smaller than the first number, and one or more readout electronics electronically coupled to the first number of PC sensors and the second number of EID detectors, the one or more readout electronics disposed on an opposite side of the second layer with respect to the incident x-rays, the one or more readout electronics receiving signals from the first number of PC sensors through a plurality of circuits wired along a plurality of paths to a plurality of cast reflectors of the second number of EID detectors or through a flexible cable around the second number of EID detectors.

[0101] When introducing elements of various embodiments of the present disclosure, the articles "a", "an", and "the" are intended to mean that there are one or more of the element. The terms "first", "second", and the like do not denote an order, quantity, or importance, but are used to distinguish one element from another. The terms "including", "comprising", and "having" are intended to be inclusive and mean that there may be additional elements other than the listed elements. As used herein, when the terms "connected", "coupled", and the like are used, one object (e.g., a material, element, structure, member, etc.) can be connected to or coupled to another object, regardless of whether the one object is directly connected or coupled to the other object, or whether there are one or more intervening objects between the one object and the other object. In addition, it should be understood that reference to "one embodiment" or "embodiments" of the present disclosure is not intended to be interpreted as excluding the existence of additional embodiments that also incorporate the referenced features.

[0102] In addition to the modifications shown above, many other modifications and alternative structures will be contemplated by those skilled in the art without departing from the spirit and scope of the present description, and the claims are intended to include such modifications and structures. Thus, while the above information has been described in particular detail with respect to what is presently believed to be the most practical and preferred embodiments, it will be apparent to those skilled in the art that many modifications can be made in form, function, operation, and use (including, but not limited to, the present invention) without departing from the principles and concepts described herein. In addition, the examples and embodiments herein are meant to be merely illustrative in all respects and should not be construed as limiting in any manner. [Explanation of symbols]

[0103] 230 Image Reconstructor 300 detector array 400 Multi-layer detector structure 402, 502 PC sensor array 404, 515, 697 EID detector arrays 406, 601 X-ray beam 414, 694 PC Sensor 416, 663, 664, 665 EID detector 440, 442 boundary 501 Incident X-ray beam 514 Wirebond 518 Connection 602 Silicon Chip 610 Sensors 614 Circuit 616, 676, 695, 696 ASICs 693 Flexible Cable 700 methods

Claims

1. A computed tomography (CT) detector system (300, 400), comprising: a layer of energy integrating detectors (EIDs) (404) disposed on an opposite side of the layer of photon counting (PC) sensors (402) with respect to the incident x-rays (406), the number of the PC sensors being greater than the number of the EID detectors; The CT detector system (300, 400) further includes an image processing unit (230) configured to correct the PC data using the EID data, denoise images reconstructed from the EID data and the PC data using a deep learning convolutional neural network (CNN) trained with a plurality of pairs of images, and increase the resolution of the reconstructed images, each pair of images including a target image reconstructed from a first signal from a layer of PC sensors (402) and an input image reconstructed from a second signal from a layer of EID detectors (404), the EID data and the PC data being acquired simultaneously from a line path of the same patient.

2. The CT detector system (300, 400) of claim 1, wherein both the PC data and the EID data are used to generate spectral and non-spectral information.

3. 2. The CT detector system of claim 1, wherein a boundary between each PC sensor of the plurality of PC sensors is not aligned with a boundary between each EID detector of the plurality of EID detectors.

4. 2. The CT detector system (300, 400) of claim 1, further comprising a readout electronic circuit section (518) arranged on the opposite side of the layer of the EID detector (515) with respect to incident X-rays (501), and a signal from each PC sensor (502) of the plurality of PC sensors is transmitted to the readout electronic circuit section (518) through a wire bond (514) wired between the EID detectors (515).

5. 2. The CT detector system (300, 400) of claim 1, further comprising a first set of readout electronics (676) and a second set of readout electronics (616), wherein a signal from each EID detector (663, 664, 665) of the plurality of EID detectors is transmitted to the first set of readout electronics (676) and a signal from each PC sensor (610) of the plurality of PC sensors is transmitted to the second set of readout electronics (616).

6. 6. The CT detector system (300, 400) of claim 5, wherein a first set of readout electronics (695) and a second set of readout electronics (696) are positioned on opposite sides of the EID detector layer (697) with respect to incident X-rays (601), and signals from each PC sensor (694) of the plurality of PC sensors are transmitted to the second set of readout electronics (696) through one or more flexible cables (693) routed around the EID detector layer (697).

7. 6. The CT detector system of claim 5, wherein a first set of readout electronics is positioned on an opposite side of the EID detector layer with respect to incident X-rays, and a second set of readout electronics is positioned on one side of the PC sensor layer.

8. 8. The CT detector system of claim 7, wherein the PC sensors (610) of the layer of PC sensors are arranged in the form of a two-dimensional PC sensor array, and each PC sensor array (651, 652, 653, 654, 655, 656) comprises a semiconductor material (602) attached to a printed circuit board (PCB) including a readout electronic circuit section (616) of a second set of readout electronic circuit sections (616) and a plurality of circuits (614), each circuit (614) of the plurality of circuits (614) electrically coupling a PC sensor (610) of the PC sensor array (651, 652, 653, 654, 655, 656) to the readout electronic circuit section (616).

9. 2. The CT detector system of claim 1, wherein the image processing unit is configured to correct the PC data using the EID data by adjusting a pile-up correction applied to photon count values ​​detected at the PC sensors of the layer of PC sensors based on both an energy spectrum of the PC data and an incident flux rate detected at an EID detector arranged on an opposite side of the layer of PC sensors relative to the incident X-rays.

Images