Active pixel sensors for photon-counting X-ray detectors

Abstract

To reduce the noise and power consumption of a system.SOLUTION: A photon-counting detector includes a plurality of detector sub-modules including semiconductor substrates. Each detector sub-module includes a plurality of active pixels configured to act as detector elements and disposed on the semiconductor substrate. Each detector sub-module further includes a plurality of traces extending from the plurality of active pixels to readout circuitry. Each active pixel of the plurality of active pixels is coupled to a respective trace of the plurality of traces. Each active pixel includes an amplification stage configured to generate an output signal based on a current pulse output generated by the active pixel. The photon counting detector also includes the readout circuitry configured to directly read out the output signals from the plurality of active pixels, where the readout circuitry is separate from the plurality of detector sub-modules.SELECTED DRAWING: Figure 5

Description

[Background technology]

[0001] The subject matter disclosed herein relates to X-ray detectors, and more particularly to active pixel sensors for photon-counting X-ray detectors.

[0002] Non-invasive imaging techniques allow images of the internal structure and features of an object (patient, manufactured item, baggage, luggage, passenger) to be obtained without physical contact.

[0003] For example, in imaging techniques using x-rays, x-ray radiation passes through an object of interest (such as a patient) and some of the x-ray radiation strikes a detector, where intensity data is collected. In digital x-ray systems, the detector generates signals that represent the amount or intensity of radiation that strikes different pixel areas on the detector surface. These signals are then processed to produce an image, which is then displayed for viewing.

[0004] In one such X-ray-based technology, known as computed tomography (CT), a scanner can project a fan- or cone-shaped X-ray beam from an X-ray source at multiple view angle locations around the object being imaged (e.g., a patient). The X-ray beam is attenuated as it traverses the object and is detected by a set of detector elements, which generate a signal representing the intensity or number of incident X-rays reaching the detector. This signal is processed to generate data representing the line integral of the linear attenuation coefficient of the object along the X-ray path. These signals are typically referred to as "projection data" or simply "projections." Reconstruction techniques (such as filtered back projection) can be used to generate images representing cross-sectional slices or three-dimensional (3D) volumes of regions of interest in the patient or imaged object. In medical applications, pathologies or other structures of interest can be located or identified from the reconstructed images or rendered volumes.

[0005] Some CT detectors contain photon-counting detectors, which convert each detected X-ray photon, measured in energy units (keV), into a voltage pulse measured in pulse height units (mV). X-ray photons are absorbed in semiconductor materials (e.g., cadmium zinc telluride (CZT), silicon, etc.), generating a photocharge proportional to the X-ray photon energy. A photodiode or diode separates the electron-hole pairs and generates a current pulse at its output. This current is fed to an application-specific integrated circuit (ASIC), which tracks the individual current pulses, determines the energy of the X-ray photons that generated them, and assigns them to the appropriate energy bin.

[0006] In state-of-the-art photon-counting detectors, the semiconductor sensor consists of an array of photodiode pixels, a separate layer to which the ASIC is directly attached by wire bonding. This heterogeneous integration creates a large capacitive load on the ASIC's input sensitive amplifier (CSA), increasing system noise and power consumption. Summary of the Invention

[0007] A summary of certain embodiments disclosed herein is provided below. It should be understood that these aspects are presented merely to provide the reader with a brief summary of certain embodiments, and that these aspects are not intended to limit the scope of the present disclosure. Indeed, the present disclosure may encompass a variety of aspects not set forth below.

[0008] In one embodiment, a photon-counting detector is provided. The photon-counting detector includes a plurality of detector sub-modules. Each detector sub-module includes a semiconductor substrate. Each detector sub-module includes a plurality of active pixels disposed on the semiconductor substrate and configured to function as detector elements. Each detector sub-module further includes a plurality of traces extending from the plurality of active pixels to a readout circuit. Each active pixel of the plurality of active pixels is coupled to a respective one of the traces. Each active pixel includes an amplifier stage configured to generate an output signal based on a current pulse output generated by the active pixel. The photon-counting detector also includes a readout circuit configured to directly read out the output signals from the plurality of active pixels, the readout circuit being separate from the plurality of detector sub-modules.

[0009] In another embodiment, a computed tomography (CT) imaging system is provided. The CT imaging system includes a photon-counting detector. The photon-counting detector includes at least one detector sub-module. The at least one detector sub-module includes a plurality of active pixels configured to function as detector elements disposed on the semiconductor substrate. The at least one detector sub-module further includes a plurality of traces extending from the plurality of active pixels to a readout circuit. Each active pixel of the plurality of active pixels is coupled to a respective one of the traces. Each active pixel includes an amplifier stage configured to generate an output signal based on a current pulse output generated by the active pixel. The photon-counting detector includes a readout circuit configured to directly read out the output signal from the plurality of active pixels, the readout circuit being separate from the at least one detector sub-module.

[0010] In a further embodiment, a photon-counting detector is provided. The photon-counting detector includes at least one detector sub-module. The at least one detector sub-module includes a semiconductor substrate. The at least one detector sub-module also includes a plurality of active pixels configured to function as detector elements disposed on the semiconductor substrate. The at least one detector sub-module includes a plurality of traces extending from the plurality of active pixels to a readout circuit. Each active pixel of the plurality of active pixels is coupled to a respective one of the traces. Each active pixel is configured to function as a passive integrator, and each active pixel includes a transconductance amplifier including a single transistor configured to generate an output signal based on a current pulse output generated by the active pixel. The photon-counting detector also includes a readout circuit configured to directly read out the output signal from the plurality of active pixels, the readout circuit being separate from the at least one detector sub-module, and the readout circuit including a current source biasing the single transistor.

[0011] In a further embodiment, a photon-counting detector is provided. The photon-counting detector includes a plurality of detector sub-modules. Each detector sub-module includes a semiconductor substrate. Each detector sub-module also includes a plurality of active pixels configured to function as detector elements disposed on the semiconductor substrate. Each detector sub-module further includes a plurality of traces extending from the plurality of active pixels to a readout circuit. Each active pixel of the plurality of active pixels is coupled to a respective one of the traces. Each active pixel includes a bipolar junction transistor configured to generate an output signal, the output signal being an amplified signal of a current pulse generated by the active pixel in response to an X-ray photon impinging on the semiconductor substrate. The photon-counting detector also includes a readout circuit configured to directly read out the output signal from the plurality of active pixels, the readout circuit being separate from the plurality of detector sub-modules. [Brief explanation of the drawings]

[0012] These and other features, aspects, and advantages of the present subject matter may be better understood from the following detailed description when read in conjunction with the accompanying drawings. [Figure 1] FIG. 1 is a block diagram of a CT system according to an aspect of the present disclosure. [Figure 2] 1 is a schematic diagram illustrating an example of modular X-ray detector sub-modules arranged side by side and stacked on top of each other, according to an aspect of the present disclosure. FIG. [Figure 3] FIG. 1 is a cross-sectional view of a portion of an X-ray detector sub-module, according to an embodiment of the present disclosure. [Figure 4] 1 is a schematic diagram showing a portion of a prior art X-ray detector sub-module having passive pixels; [Figure 5] FIG. 2 is a schematic diagram illustrating a portion of an X-ray detector sub-module having active pixels, according to an aspect of the present disclosure. [Figure 6]1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having active pixels (e.g., passive integrators) according to an embodiment of the present disclosure. [Figure 7] 1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having active pixels (eg, CSAs integrated into the pixels) according to an embodiment of the present disclosure. [Figure 8] 1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having multiple active pixels (e.g., passive integrators) according to an embodiment of the present disclosure. [Figure 9] 1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having multiple active pixels (e.g., each pixel having an integrated CSA), according to an embodiment of the present disclosure. [Figure 10] 1 is a flowchart of a method for manufacturing an X-ray detector sub-module having a plurality of pixels, according to an aspect of the present disclosure. [Figure 11] 1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having active pixels (e.g., pixels having PNP-type bipolar junction transistors) according to an embodiment of the present disclosure. [Figure 12] 1 is a schematic diagram illustrating a portion of an X-ray detector sub-module having active pixels (e.g., pixels having NPN-type bipolar junction transistors) according to an embodiment of the present disclosure. [Figure 13] 1 is a cross-sectional view of a portion of an X-ray detector sub-module having a bipolar junction transistor (e.g., NPN type) illustrating its operation, according to an embodiment of the present disclosure. [Figure 14] 1 is a schematic diagram of a portion of an X-ray detector sub-module having bipolar junction transistors (e.g., having an NPN type), according to an embodiment of the present disclosure. [Figure 15] 1 is a schematic diagram of a portion of an X-ray detector sub-module having bipolar junction transistors (e.g., having a PNP type), according to an embodiment of the present disclosure. DETAILED DESCRIPTION OF THE INVENTION

[0013] Described below are one or more specific embodiments. As a result of the intent to provide a concise description of these embodiments, not all features of an actual implementation are described herein. It should be understood that the development of any actual implementation, like any engineering or design project, requires numerous implementation-specific decisions to achieve the developer's particular goals (including compliance with system-related and business-related constraints that may vary from implementation to implementation). Moreover, it should be understood that such a development effort might be complex and time-consuming, but would nevertheless be a routine undertaking of design, fabrication, and manufacture for those of ordinary skill in the art having the benefit of this disclosure.

[0014] When introducing elements of various embodiments of the present subject matter, the articles "a," "an," "the," and "said" are intended to mean that there are one or more of the element. The terms "comprising," "including," and "having" are intended to be inclusive and mean that there may be additional elements other than the listed elements. Moreover, the numerical examples in the following description are intended to be non-limiting, and therefore, additional numbers, ranges, and percentages are within the scope of the disclosed embodiments.

[0015] While the following discussion is described in the context of medical imaging, it should be understood that the present technology is not limited to such a medical context. Indeed, the examples are shown and described in such a medical context merely to facilitate explanation by providing examples of real-world implementations and applications. However, the present techniques can also be utilized in other contexts, such as non-destructive testing of manufactured parts or goods (i.e., quality control or quality screening applications) and / or non-invasive testing of packages, boxes, luggage, etc. (i.e., security or screening applications). In general, the present techniques are useful in imaging or screening contexts where photon-counting detectors are used.

[0016] Energy-resolving photon-counting detectors can provide spectral information not available from conventional energy-integrating detectors. One type of energy-discriminating photon-counting detection technology employs silicon strips as the direct-conversion sensor material. Using silicon as the direct-conversion material allows for higher counts than those achieved using other direct-conversion materials (e.g., CZT or CdTe). In certain embodiments, the detector can be positioned edge-on, allowing for the absorption depth to be selected to any length, thereby increasing absorption efficiency and eliminating the need for very high voltages to fully deplete the detector. The detector elements of a detector submodule or sensor (especially those along the edge of the detector submodule or sensor) are typically coupled to an ASIC (separate from the detector submodule) via direct wire bonding. This heterogeneous integration introduces a significant capacitive load to the input CSA of the ASIC, increasing system noise and power consumption.

[0017] The present approach alleviates this problem by implementing a practical active pixel sensor for a photon counting system (photon counting detector). In particular, each detector sub-module includes one or more active pixels (e.g., diodes and associated circuitry) and an amplification stage to overcome the effects of detector capacitance noise. The amplification stage is integrated into the pixel and provides current gain to increase the signal-to-noise ratio and suppress the effects of detector capacitance noise. The output signal from the amplification stage is read out directly by a readout circuit (e.g., an ASIC) separate from the detector sub-module. In a specific embodiment, each active pixel is configured to function as a passive integrator. In this embodiment, the amplification stage includes a transconductance amplifier having a single transistor. In this embodiment, a current source is disposed in the readout circuit and biases the single transistor. In a specific embodiment, the amplification stage includes a charge-sensitive amplifier. In this embodiment, the amplification stage includes a first transistor and a second transistor, where the first transistor is a charge-sensitive amplifier and the second transistor is configured to convert a voltage output into a current. In this embodiment, a current source is integrated into each active pixel, and the active pixels are free of bias circuitry. These techniques improve the performance of silicon-based photon-counting detectors (such as computed tomography detectors or other suitable types of radiography detectors). In particular, the disclosed techniques directly improve image quality by reducing the effects of electronic noise (by improving the signal-to-noise ratio). Furthermore, the disclosed techniques reduce both system complexity and cost due to low power consumption of the readout circuitry (e.g., ASIC).

[0018] With the foregoing in mind, FIG. 1 illustrates an embodiment of an imaging system 10 for acquiring and processing image data utilizing the active pixels described herein. While the following embodiments are described with respect to a computed tomography (CT) imaging system, these embodiments may also be utilized with other imaging systems (e.g., X-ray, PET, CT / PET, SPECT, nuclear CT, etc.). In the illustrated embodiment, system 10 is a computed tomography (CT) system designed to acquire X-ray projection data, reconstruct the projection data to generate images, and process the image data for display and analysis. CT imaging system 10 includes one or more X-ray sources 12 (e.g., one or more X-ray tubes or solid-state emitting structures capable of generating X-rays at one or more locations and / or at one or more energy spectra during an imaging session).

[0019] In certain embodiments, the source 12 may be positioned near a collimator 22 used to define the size and shape of one or more x-ray beams 20 passing through a region where a subject 24 (e.g., a patient) or object of interest is located. At least a portion of the x-rays are attenuated by the subject 24. The attenuated x-rays 26 impinge on a detector array 28 formed by a plurality of detector elements (e.g., pixels). As described herein, the detector 28 may be a photon-counting detector, including an energy-discriminating photon-counting detector, the output of which conveys information regarding the number and energy of photons impinging on the detector over a time period corresponding to a scan or imaging session and at a measurement location. In certain such embodiments, the energy-discriminating photon-counting detector may be a direct conversion detector (i.e., without a scintillator means), such as a silicon strip-based detector. In certain embodiments, the detector array 28 may be formed by a plurality of detector sub-modules or sensors, each having a plurality of detector elements, such as photodiodes or diodes. In certain embodiments, detector array 28 and detector sub-modules may be edge-on detectors and edge-on detector sub-modules configured for edge-illumination of x-rays (i.e., x-rays passing through the edge of the detector sub-module). In particular, detector array 28 may be similar in structure to the detectors disclosed in U.S. Patent Application Publication No. 2019 / 0383955, filed February 19, 2019, and entitled "X-ray Detector System Design," which is incorporated herein in its entirety for all purposes.

[0020] Each detector element produces an electrical signal that represents the intensity (e.g., energy and number of incident X-ray photons) at the position of the detector element when the beam strikes the detector 28. The electrical signals are acquired and processed to produce one or more scan data sets.

[0021] The system controller 30 commands the operation of the imaging system 10 to execute inspection and / or calibration protocols and to process acquired data. With respect to the x-ray source 12, the system controller 30 provides power, focal spot position, control signals, etc. for x-ray inspection sequences. The detector 28 is coupled to the system controller 30, which commands the acquisition of signals generated by the detector 28. Additionally, the system controller 30, through a motor controller 36, may control the operation of a linear positioning subsystem 32 and / or a rotational subsystem 34 used to move components of the imaging system 10 and / or move the subject 24. The system controller 30 may include signal processing circuitry and associated memory circuitry. In such an embodiment, the memory circuitry may store programs, routines, and / or coded algorithms executed by the system controller 30 to operate the imaging system 10 (including the x-ray source 12) and process data acquired by the detector 28 according to the steps and processes described herein. In one embodiment, the system controller 30 may be implemented as all or part of a processor-based system (e.g., a general-purpose computer system or a special-purpose computer system).

[0022] The sources 12 may be controlled by an X-ray controller 38 included in the system controller 30. The X-ray controller 38 may be configured to provide power and timing signals to the sources 12. Additionally, in some embodiments, the X-ray controller 38 may be configured to selectively operate the sources 12 such that tubes or emitters at different locations within the system 10 operate synchronously with or independently of one another.

[0023] The system controller 30 may include a data acquisition system (DAS) 40. The DAS 40 receives data (e.g., sampled analog signals from the detector 28) collected by the readout electronics (e.g., ASIC) of the detector 28. The DAS 40 may then convert the data to digital signals for subsequent processing by a processor-based system (e.g., a computer 42). In other embodiments, the detector 28 may convert the sampled analog signals to digital signals before transmitting them to the data acquisition system 40. The computer 42 may include processing circuitry 44 (e.g., image processing circuitry). The computer 42 may include or communicate with one or more non-transitory memory devices 46 that may store data to be processed by the computer 42, data to be processed by the computer 42, or instructions to be executed by a processor (e.g., the processing circuitry 44) of the computer 42. For example, processing circuitry 44 of computer 42 may execute one or more sets of instructions stored in memory 46, which may be memory of computer 42, memory of a processor, firmware, or similar implementations.

[0024] The computer 42 may also be adapted to control functions implemented by the system controller 30 (i.e., scanning operations and data acquisition) in response to commands and scanning parameters provided by an operator, for example, through an operator workstation 48. The system 10 may also include a display 50 coupled to the operator workstation 48, which allows the operator to view relevant system data, imaging parameters, raw imaging data, reconstructed data, etc. The system 10 may also include a printer 52 coupled to the operator workstation 48, which is configured to print any desired measurements. The display 50 and printer 52 may be connected directly to the computer 42 or through the operator workstation 48. The operator workstation 48 may also include or be coupled to a picture archiving and communication system (PACS) 54. The PACS 54 may be coupled to a remote system 56, a radiology information system (RIS), a hospital information system (HIS), or an internal or external network, allowing others at different locations to access the image data.

[0025] FIG. 2 is a schematic diagram illustrating an example of modular X-ray detector submodules 58 (e.g., detector sensors) arranged side by side and stacked. The detector submodules 58 may be edge-on detector submodules. As shown, X-rays enter the detector submodules 58 from an edge 59. A guard ring may extend along the edge 59 of the detector submodules 58 to protect the detector submodules 58 from electrical breakdown and to insulate the detector area from excessive leakage current. In certain embodiments, the detector submodules 58 may be planar modules. The X-ray detector submodules 58 may be stacked on top of each other to form larger detector modules, and the detector modules may be assembled side by side to build an overall X-ray detector. In general, multiple detector submodules 58 may be arranged side by side, e.g., in a slightly curved overall configuration substantially perpendicular to the z-direction. In certain embodiments, multiple detector submodules 58 may be stacked one on top of the other in the z-direction.

[0026] As shown, each detector sub-module 58 includes multiple detector elements 60 (e.g., pixels, such as photodiodes or diodes). The detector elements 60 may be elongated electrodes (e.g., metal electrodes of a photodiode) extending toward the focal point of an x-ray system. Depending on the detector topology, the detector elements 60 may correspond to pixels. In certain embodiments, the detector sub-module 58 may be a depth-segmented detector sub-module having multiple detector strips 62, each having multiple depth segments 65. As shown, each strip 62 includes a first segment 64, a second segment 66, and a third segment 68 at different depths along the detection line (relative to the focal point). As shown, at least a portion of each of the segments 64, 66, and 68 are collinear. The number of segments 65 may be different (e.g., one to three or more). In such a depth segment detector sub-module 58, each depth segment 65 can be considered an individual detector element (where each depth segment is associated with its own individual charge collection electrode). In certain embodiments, the circuitry can logically treat the depth segments 65 of a single strip 62 as a single detector element.

[0027] The detector sub-modules 58 may have other shapes. In certain embodiments, the detector sub-modules 58 may have a parallelogram shape, a trapezoid shape, a triangular shape, or another shape. In certain embodiments, one or more edges 59 of the detector sub-modules 58 may be beveled. The detector elements 60 may have other shapes. In certain embodiments, the detector elements 60 disposed along the beveled side edges 69 of the detector sub-modules 58 may include tapered edge segments (e.g., trapezoidal or triangular segments and / or truncated trapezoidal or triangular segments with rounded corners). In certain embodiments, the segments 65 of the strips 62 closest to the beveled side edges of the detector sub-modules 58 may be oriented to extend into the regions of the adjacent strips 62. In certain embodiments, the segments 65 may be formed at an angle.

[0028] FIG. 3 is a cross-sectional view of a portion of an X-ray detector submodule 70. The X-ray detector submodule 70 includes a semiconductor layer 72. The semiconductor layer 72 is made of silicon. In particular embodiments, the semiconductor layer 72 may be made of gallium arsenide, zinc cadmium telluride, or other semiconductor materials. Detector elements or segments 74, 76 (e.g., metal electrodes of photodiodes) are disposed on the semiconductor layer 72. The electrodes may be made of aluminum. In particular, the electrodes 74, 76 are disposed on doped implants 78, 80 (e.g., p-type or n-type silicon implants, depending on whether the silicon in the semiconductor layer is n-type or p-type) disposed on the semiconductor layer 72. The detector elements 74, 76 may be segments of different strips disposed adjacent to each other. The detector elements 74, 76 and the doped implants 78, 80 are spaced apart in the semiconductor layer 72, forming a gap 82 between them. The X-ray detector submodule 70 includes an electrical insulation layer 84 extending between adjacent electrodes 74, 76. The electrical insulation layer 84 can be silicon dioxide, silicon nitride, polyimide, spin-on glass, or other insulating material. One or more wiring traces 86 (e.g., metal traces) are routed in the gap 82 between the electrodes 74, 76. As shown, the wiring traces 86 are disposed in the electrical insulation layer 84. As shown, the wiring traces 86 are evenly spaced across the gap 82. In certain embodiments, the wiring traces 86 can be routed as close as possible to the edges of the electrodes 74, 76. The wiring traces 86 may be coupled to the electrodes 74, 76 or to different electrodes. The wiring traces 86 are routed along the gap 82 (and possibly other gaps) to readout circuitry. As shown, a passivation layer 88 is disposed over these components of the X-ray detector submodule 70. The passivation layer 88 can be formed of silicon oxide, silicon nitride, or other insulator.

[0029] FIG. 4 is a schematic diagram illustrating a portion of a prior art X-ray detector sub-module 90 (detector sensor) having passive pixels (i.e., instantaneous readout without pixel-by-pixel amplification). As shown, the X-ray detector sub-module 90 includes a photodiode or diode 92 (e.g., pixel or detector element). The X-ray detector sub-module 90 may include multiple photodiodes 92. A voltage bias 94 is applied to the photodiode 92. The photodiode 92 is coupled to a trace 96 (e.g., a conductive metal trace). The trace 96 is a data line (e.g., a dedicated data line) coupled to a readout circuit 98 (e.g., a front end of an ASIC) for reading out a signal (e.g., a current pulse) output from the photodiode 92 in response to detection of an incident X-ray photon (e.g., a photon absorbed by a semiconductor layer). The readout circuit 98 is a separate circuit from the X-ray detector sub-module 90 (i.e., the readout circuit 98 is not located in the semiconductor layers of the X-ray detector sub-module 90). Each photodiode 92 is coupled to a dedicated channel (e.g., a dedicated readout channel) of a readout circuit 98. The traces 96 of the photodiodes 92 are coupled to the readout circuit 98 by direct wire bonding. This heterogeneous integration presents a large capacitive load to the input CSA 100 of the readout circuit 98 (e.g., an ASIC). In particular, noise 102 on the traces 96 or on the readout circuit 98 is dissipated by the detector capacitance (C det ) 104 and the feedback capacitance (C f ) 106. In addition to increasing noise, a large capacitive load also increases the power consumption of the system.

[0030] FIG. 5 is a schematic diagram illustrating a portion of an X-ray detector sub-module 108 (detector sensor) having active pixels (i.e., pixels that include some form of amplification). The X-ray detector sub-module 108 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 110 includes a photodiode or diode 112 (e.g., an active pixel or detector element). The X-ray detector sub-module 110 may include multiple photodiodes 112. A voltage bias 114 is applied to the photodiode 112. The photodiode 112 is coupled to an amplifier stage 116 (i.e., the amplifier stage 116 is integrated into the active pixel). The amplifier stage 116 is configured to generate an output signal based on a current pulse output generated by the active pixel in response to detection of an incident X-ray photon (e.g., absorbed by a semiconductor layer). The amplifier stage 116 is coupled to a trace 118 (e.g., a conductive metal trace). The trace 118 is a data line (e.g., a dedicated data line) coupled to a readout circuit 120 (e.g., the front end of an ASIC) for reading out the signal output from the amplifier stage 116. The readout circuit 120 is separate from the X-ray detector submodule 108 (i.e., not located in a semiconductor layer of the X-ray detector submodule 108). Each photodiode 116 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit 120. The trace 118 of the photodiode 112 is directly coupled to the readout circuit 120 by direct wire bonding. The photodiode 112 does not accumulate detected charge and is therefore directly coupled to the readout circuit 120, which, in certain embodiments, is provided to an input CSA 122 of the readout circuit 120 (e.g., an ASIC). Any noise 124 from the trace 118 or the readout circuit 120 is still present. However, the amplifier stage 116 boosts the signal from the photodiode 112 (i.e., provides current gain) and increases the signal-to-noise ratio. Furthermore, the detector capacitance (C det ) 125 (i.e., trace capacitance and parasitic capacitance) on noise 124 is reduced.

[0031] In certain embodiments, each active pixel (e.g., photodiode 112) is configured to function as a passive integrator. When each active pixel is a passive integrator, amplification stage 116 includes a transconductance amplifier having a single transistor. Also, when each active pixel is a passive integrator, a current source is disposed in readout circuit 120 to bias the single transistor. In certain embodiments, amplification stage 116 includes a charge-sensitive amplifier. When amplification stage 116 includes a charge-sensitive amplifier, amplification stage 116 includes a first transistor and a second transistor, where the first transistor is a charge-sensitive amplifier and the second transistor is configured to convert a voltage output to a current. Also, when amplification stage 116 includes a charge-sensitive amplifier, a current source is integrated into each active pixel.

[0032] 5, the image data obtained from the photon-counting detector is of improved quality and the power consumption of the readout circuit 120 is reduced.

[0033] In certain embodiments, instead of one or more diodes, the X-ray detector sub-module 110 includes multiple bipolar junction transistors (e.g., vertical bipolar junction transistors) as active pixels (see FIGS. 11-15). In these embodiments, a separate amplification stage is not required. Instead, the bipolar junction transistors amplify the pulse signals generated by X-ray photons impinging on the semiconductor layers of the X-ray detector sub-module 110.

[0034] FIG. 6 is a schematic diagram illustrating a portion of an X-ray detector sub-module 126 (detector sensor) having active pixels (e.g., passive integrators). The X-ray detector sub-module 126 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 126 includes a photodiode or diode 128 (e.g., an active pixel or detector element). The X-ray detector sub-module 126 may include multiple photodiodes 128 (see FIG. 8). A voltage bias 130 is applied to the photodiode 128. The photodiode 128 is coupled to an amplifier stage 132 (i.e., the amplifier stage 132 is integrated into the active pixel). The amplifier stage 132 is configured to generate an output signal based on a current pulse output generated by the active pixel in response to detection of an incident X-ray photon (e.g., an incident X-ray photon absorbed by a semiconductor layer).

[0035] Each active pixel (e.g., photodiode 128) is configured to function as a passive integrator. Amplification stage 132 includes a transconductance amplifier 134 (M1). In a particular embodiment, transconductance amplifier 134 is a p-channel metal-oxide semiconductor (PMOS). Transconductance amplifier 134 includes a single transistor 136. Transconductance amplifier 134 has a gain (g m ) from the photodiode 128. pix ) to the current (I out ) The amplifier stage 132 includes a bias and reset circuit 138 coupled to the transconductance amplifier 134. The amplifier stage 132 converts the pixel capacitor 140 (C pix) and is configured to discharge the integrator (i.e., input gate capacitance). The bias reset circuit 138 is also configured to provide a reset path for the pixel capacitor 140. The bias reset circuit 138 is configured so that only an output signal (not a bias or DC current) based on the current pulse output generated by the active pixel resets the feedback capacitor (C f (see

[0036] The amplifier stage 132 is coupled to a trace 142 (e.g., a conductive metal trace). The trace 142 is a data line (e.g., a dedicated data line) coupled to a readout circuit 144 (e.g., the front end of an ASIC) for reading out the signal output from the amplifier stage 132. A current source 146 is disposed in the readout circuit 144, and the current source 146 biases the single transistor 136. The readout circuit 144 is separate from the X-ray detector submodule 126 (i.e., not disposed in the semiconductor layer of the X-ray detector submodule 126). Each photodiode 128 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit 144. The trace 142 of the photodiode 128 is directly coupled to the readout circuit 144 by direct wire bonding. The photodiodes 128 do not accumulate detected charge and are therefore directly coupled to the readout circuit 144. The amplifier stage 132 boosts the signal from the photodiode 128 (i.e., provides current gain) and increases the signal-to-noise ratio. det ) 148 is suppressed from affecting noise.

[0037] 6, the image data obtained from the photon-counting detector is of improved quality and the power consumption of the readout circuit 144 is reduced.

[0038] FIG. 7 is a schematic diagram illustrating a portion of an X-ray detector sub-module 150 (detector sensor) having active pixels (e.g., CSAs integrated into the pixels). The X-ray detector sub-module 150 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 150 includes a photodiode or diode 152 (e.g., an active pixel or detector element). The X-ray detector sub-module 126 may include multiple photodiodes 152 (see FIG. 9). A voltage bias 154 is applied to the photodiode 152. The photodiode 152 is coupled to an amplifier stage 156 (i.e., the amplifier stage 156 is integrated into the active pixel). The amplifier stage 156 is configured to generate an output signal based on a current pulse output generated by the active pixel in response to detecting an incident X-ray photon (e.g., an incident X-ray photon absorbed by a semiconductor layer).

[0039] Each active pixel (e.g., photodiode 152) is configured to function as a passive integrator. Amplification stage 156 includes a charge-sensitive amplifier 158 (M1). The charge-sensitive amplifier 158 is integrated with the active pixel. Amplification stage 156 includes a first transistor 160 and a second transistor 161. In a particular embodiment, both first transistor 160 and second transistor 161 are PMOS. In a particular embodiment, first transistor 160 is a PMOS, and second transistor 161 is an n-channel metal-oxide semiconductor (NMOS). No biasing of first transistor 160 is required. If second transistor 161 is an NMOS, second transistor 161 functions as a source follower (e.g., a voltage buffer). As a result, readout circuit 170 (e.g., an ASIC) only needs a shaper and a discriminator. If the second transistor 161 is a PMOS, the second transistor 161 is a two-stage gain and the readout circuit 170 requires a CSA, a shaper, and a discriminator. The first transistor 160 is a charge sensitive amplifier 158. The CSA amplifier 158 is a transconductance amplifier that provides gain. The second transistor 161 (M2) converts the voltage from the charge sensitive amplifier 158 into a current (I out ) and a second transistor 161 provides additional gain. As shown in FIG. 7, the amplifier stage 156 includes a current source 162 integrated into each active pixel. The amplifier stage 156 includes a resistor (R) 164. The amplifier stage 156 is connected to a feedback capacitor 166 (C f ) is also included.

[0040] The amplifier stage 156 is coupled to a trace 168 (e.g., a conductive metal trace). The trace 168 is a data line (e.g., a dedicated data line) coupled to a readout circuit 170 (e.g., the front end of an ASIC) for reading out the signal output from the amplifier stage 156. The readout circuit 170 is separate from the X-ray detector submodule 150 (i.e., the readout circuit 170 is not disposed in a semiconductor layer of the X-ray detector submodule 150). Each photodiode 152 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit 170. The trace 168 of the photodiode 152 is directly coupled to the readout circuit 170 by direct wire bonding. The photodiode 152 does not accumulate detected charge and is therefore directly coupled to the readout circuit 170. The amplifier stage 156 boosts the signal from the photodiode 152 (i.e., provides current gain), increasing the signal-to-noise ratio. Additionally, the detector capacitance (C det ) 172 is less susceptible to noise.

[0041] 7, the image data obtained from the photon-counting detector is of improved quality, and the power consumption of the readout circuit 170 is reduced.

[0042] FIG. 8 is a schematic diagram illustrating a portion of an X-ray detector sub-module 174 (detector sensor) having multiple active pixels (e.g., passive integrators). The X-ray detector sub-module 174 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 174 includes multiple photodiodes or diodes 176 (e.g., active pixels or detector elements). A voltage bias 178 is applied to each photodiode 176. Each photodiode 176 is coupled to a respective amplifier stage 180 (i.e., each amplifier stage 180 is integrated with a respective active pixel). Each amplifier stage 180 is configured to generate an output signal based on a current pulse output generated by a respective active pixel in response to detection of an incident X-ray photon (e.g., an incident X-ray photon absorbed by a semiconductor layer).

[0043] Each active pixel (e.g., photodiode 176) is configured to function as a passive integrator. Each amplification stage 180 includes a transconductance amplifier 182 (M1). Each transconductance amplifier 182 includes a single transistor 184. Each transconductance amplifier 182 has a gain (g m ) from its respective photodiode 176. In a particular embodiment, each transconductance amplifier 182 is a PMOS. Each amplifier stage 180 is a post-integration amplifier that provides the voltage (V pix ) to the current (I out ) to a pixel capacitor 188 (C pix). Each bias reset circuit 186 is configured to discharge an integrator (i.e., input gate capacitance). Each bias reset circuit 186 is also configured to provide a reset path for a pixel capacitor 188. Each bias reset circuit 186 is configured so that only an output signal (not a bias or DC current) based on the current pulse output generated by the active pixel resets the feedback capacitor (C f (see

[0044] Each amplifier stage 180 is coupled to a respective trace 190 (e.g., a conductive metal trace). Each trace 190 is a data line (e.g., a dedicated data line) coupled to a readout circuit 192 (e.g., the front end of an ASIC) for reading out a signal output from each amplifier stage 180. A respective current source 193 is disposed in the readout circuit 192 and biases a respective transistor 184. The readout circuit 190 is separate from the X-ray detector submodule 174 (i.e., not disposed in a semiconductor layer of the X-ray detector submodule 174). Each photodiode 176 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit 192. In certain embodiments, the readout channels of multiple photodiodes 176 are disposed in the same ASIC module. In certain embodiments, the readout channels of multiple photodiodes 176 are disposed in different ASIC modules. The trace 190 of the photodiode 176 is directly coupled to the readout circuit 192 by direct wire bonding. The photodiodes 176 do not accumulate detected charge and are therefore directly coupled to the readout circuitry 192. Each amplifier stage 180 boosts the signal from its respective photodiode 176 (i.e., provides current gain) and increases the signal-to-noise ratio. Additionally, the detector capacitance (C det )194 is less susceptible to noise.

[0045] 8, the image data obtained from the photon-counting detector is of improved quality, and the power consumption of the readout circuit 192 is reduced.

[0046] FIG. 9 is a schematic diagram illustrating a portion of an X-ray detector sub-module 196 (detector sensor) having multiple active pixels (e.g., each pixel integrated with a CSA). The X-ray detector sub-module 196 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 196 includes multiple photodiodes or diodes 198 (e.g., active pixels or detector elements). A voltage bias 200 is applied to each photodiode 198. Each photodiode 198 is coupled to a respective amplifier stage 202 (i.e., each amplifier stage 202 is integrated with a respective active pixel). Each amplifier stage 202 is configured to generate an output signal based on a current pulse output generated by a respective active pixel in response to detection of an incident X-ray photon (e.g., an incident X-ray photon absorbed by a semiconductor layer).

[0047] Each active pixel (e.g., photodiode 198) is configured to function as a passive integrator. Each amplifier stage 202 includes a respective charge-sensitive amplifier 204 (M1). Each charge-sensitive amplifier 204 is integrated with its active pixel. Each amplifier stage 202 includes a respective first transistor 206 and a respective second transistor 208. In a particular embodiment, both the first transistor 206 and the second transistor 208 are PMOS. In a particular embodiment, the first transistor 206 is PMOS and the second transistor 208 is NMOS. No biasing of the first transistor 206 is required. Each first transistor 206 is a charge-sensitive amplifier 204. Each CSA amplifier 204 is a transconductance amplifier that provides gain. Each second transistor 208 (M2) converts the voltage (~V) from the respective charge-sensitive amplifier 204 to a pix ) to the current (Iout ) and each second transistor 208 provides additional gain. As shown in FIG. 9, each amplifier stage 202 includes a respective current source 210 integrated into each active pixel. Each amplifier stage 202 includes a respective resistor (R) 212. Each amplifier stage 202 includes a respective feedback capacitor 214 (C f ) is also included.

[0048] Each amplifier stage 202 is coupled to a respective trace 216 (e.g., a conductive metal trace). Each trace 216 is a data line (e.g., a dedicated data line) coupled to a readout circuit 218 (e.g., the front end of an ASIC) for reading out the signal output from the respective amplifier stage 202. The readout circuit 218 is separate from the X-ray detector submodule 196 (i.e., not located in a semiconductor layer of the X-ray detector submodule 196). Each photodiode 198 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit 218. In certain embodiments, the readout channels of multiple photodiodes 198 are located in the same ASIC module. In certain embodiments, the readout channels of multiple photodiodes 198 are located in different ASIC modules. The trace 216 of each photodiode 198 is directly coupled to the readout circuit 218 by direct wire bonding. The photodiodes 198 do not accumulate detected charge and are therefore directly coupled to the readout circuit 218. Each amplifier stage 202 boosts the signal from its respective photodiode 198 (i.e., provides current gain) and increases the signal-to-noise ratio. det ) 220 is suppressed from being affected by noise.

[0049] 9, the image data obtained from the photon-counting detector is of improved quality and the power consumption of the readout circuit 218 is reduced.

[0050] 10 is a flowchart of a method 222 for fabricating an X-ray detector submodule having a plurality of active pixels (e.g., detector submodules 58 and 70 of FIGS. 2 and 3). The method 222 includes providing a semiconductor layer (block 224). The semiconductor layer is made of silicon. In certain embodiments, the semiconductor layer may be made of gallium arsenide, cadmium zinc telluride, or another semiconductor material.

[0051] The method 222 also includes disposing a plurality of active pixels and associated circuitry in the semiconductor layer (block 226). Each active pixel includes a photodiode coupled to an amplification stage configured to increase the gain of a signal output by the photodiode. In a particular embodiment, each active pixel is configured to function as a passive integrator. In this embodiment, the amplification stage includes a transconductance amplifier having a single transistor. In this embodiment, a current source is disposed in the readout circuitry and biases the single transistor. In a particular embodiment, the amplification stage includes a charge-sensitive amplifier. In this embodiment, the amplification stage includes a first transistor and a second transistor, the first transistor being a charge-sensitive amplifier and the second transistor configured to convert a voltage output to a current. In this embodiment, a current source is integrated into each active pixel.

[0052] The method 222 further includes coupling each active pixel of the plurality of active pixels to a readout circuit (block 228). The readout circuit may include one or more ASIC modules. Each amplifier stage of each active pixel is coupled to a respective trace that functions as a data line (e.g., a dedicated data line) that is coupled to a readout circuit (e.g., a front end of an ASIC) for reading out a signal output from the respective amplifier stage. The readout circuit is separate from the X-ray detector sub-module (i.e., not located in a semiconductor layer of the X-ray detector sub-module). Each photodiode is coupled (through a trace coupled to the amplifier stage) to a dedicated channel (e.g., a dedicated readout channel) of the readout circuit (e.g., an ASIC). In certain embodiments, the readout channels of the multiple photodiodes are located in the same ASIC module. In certain embodiments, the readout channels of the multiple photodiodes are located in different ASIC modules. Each trace of the photodiode is directly coupled to the readout circuit by direct wire bonding.

[0053] FIG. 11 is a schematic diagram illustrating a portion of an X-ray detector sub-module 230 (e.g., a detector sensor) having active pixels (e.g., pixels having amplifying bipolar junction transistors). The X-ray detector sub-module 230 is similar to the detector sub-modules 58 and 70 of FIGS. 2 and 3. As shown, the X-ray detector sub-module 230 includes bipolar junction transistors 232 (e.g., active pixels or detector elements). In certain embodiments (e.g., embodiments such as those shown in FIG. 12 and corresponding to the bipolar junction transistors 246 and 272 of FIGS. 13 and 14), the bipolar junction transistors 232 (e.g., vertical bipolar junction transistors) are NPN-type bipolar junction transistors. In certain embodiments (e.g., embodiments such as those shown in FIG. 11 and corresponding to the bipolar junction transistor 294 of FIG. 15), the bipolar junction transistors 232 are PNP-type bipolar junction transistors. The X-ray detector sub-module 230 may include multiple bipolar junction transistors 232. Instead of having a separate amplification stage, the amplification is performed by bipolar junction transistors 232. The bipolar junction transistors 232 are configured to generate output signals, which are amplified signals of current pulses generated by active pixels in response to detecting incident X-ray photons (e.g., incident X-ray photons absorbed by a semiconductor layer). A voltage bias 233 is applied to the bipolar junction transistors 232. The bipolar junction transistors 232 are coupled to traces 234 (e.g., conductive metal traces). The traces 234 are data lines (e.g., dedicated data lines) coupled to readout circuits 236 (e.g., the front end of an ASIC) for reading out signals output from the bipolar junction transistors 232. The readout circuits 236 are separate from the X-ray detector sub-module 230 (i.e., not located in the semiconductor layers of the X-ray detector sub-module 230). Each bipolar junction transistor 230 is coupled to a dedicated channel (e.g., a dedicated readout channel) of the readout circuits 236.The trace 234 of the bipolar junction transistor 232 is directly coupled to a readout circuit 236 by direct wire bonding. The bipolar junction transistor 232 does not store the detected charge and is therefore directly coupled to the readout circuit 236 and, in certain embodiments, is provided to an input CSA 238 (e.g., an ASIC) of the readout circuit 236. Noise 240 from the trace 234 or the readout circuit 236 is still present. However, the bipolar junction transistor 232 boosts the signal (i.e., provides current gain), increasing the signal-to-noise ratio. Additionally, the detector capacitance (C det ) 242 (i.e., trace capacitance and parasitic capacitance) contributes less to noise 240.

[0054] 13 is a cross-sectional view of a portion of the X-ray detector submodule 244 having a bipolar junction transistor 246 (e.g., an NPN-type vertical bipolar junction transistor), illustrating its operation. The X-ray detector submodule 244 includes a first surface 248 (e.g., a front surface) and a second surface 250 (rear surface) opposite the first surface 248. As shown, the bipolar junction transistor 246 includes an emitter 252 formed from a first N-type semiconductor layer 254 (e.g., N+), a base 256 formed from a P-type semiconductor layer 258 (e.g., P), and a collector 260 formed from a second N-type semiconductor layer 262 (N−). The bipolar junction transistor 246 also includes a third N-type semiconductor layer 264 (e.g., N+). The collector 260 is disposed between the base 256 and the third N-type semiconductor layer 264. Both the emitter 252 and a portion of the base 256 are located on the first surface 248, and the third N-type semiconductor layer 264 is located on the second surface 250. The third N-type semiconductor layer 264 is more heavily doped than the collector 260. The base 256 is 10 to 100 times less doped than the emitter 252, which corresponds approximately to the gain of the bipolar junction transistor 246. The P-type region of the base 256 is sufficiently highly doped and sufficiently thick, but not too highly doped or too thick. In particular, the base 256 is sufficiently doped and sufficiently thick to avoid punch-through.

[0055] The junction 266 between the base 256 and the collector 260 is reverse biased to fully deplete. The emitter 252 is grounded. The bipolar junction transistor 246 functions as a detector element or active pixel. In response to X-ray photons passing through the X-ray detector submodule 244, scattering valence electrons in a region and exciting them into the conduction band, generating electron-hole pairs (i.e., a current pulse). The holes are injected into the base 256. When an excess of holes accumulates, a forward bias is applied to the junction 268 between the emitter 252 and the base 256. Amplified electrons are injected into the base 256. In certain embodiments, the base 256 may require a bias. As a result, the bipolar junction transistor 246 generates an output signal, which is an amplified version of the current pulse generated by the active pixel in response to X-ray photons traveling through the semiconductor substrate.

[0056] 14 is a schematic diagram showing a portion of an X-ray detector submodule 270 having a bipolar junction transistor 272 (e.g., an NPN-type bipolar junction transistor). The X-ray detector submodule 270 includes a first surface 274 (front surface) and a second surface 276 (back surface) opposite the first surface 274. As shown, the bipolar junction transistor 272 includes an emitter 278 formed from a first N-type semiconductor layer 280 (e.g., N), a base 282 formed from a P-type semiconductor layer 284 (e.g., P-), and a collector 286 formed from a second N-type semiconductor layer 288 (N). The bipolar junction transistor 246 also includes a second P-type semiconductor layer 290 (e.g., P). The second P-type semiconductor layer 290 is disposed between the base 282 and the emitter 278. The second P-type semiconductor layer 290 is more heavily doped than the base 282. Both the emitter 278 and a portion of the second P-type semiconductor layer 290 are located on the first surface 274, and the collector 286 is located on the second surface 276. The P-type region is heavily doped and sufficiently thick, but not too heavily doped or too thick. Similarly to the description of FIG. 13, the bipolar junction transistor 272 generates an output signal. The output signal is an amplified version of the current pulse generated by the active pixel in response to an X-ray photon traveling through the semiconductor substrate.

[0057] 15 is a schematic diagram illustrating a portion of an X-ray detector sub-module 292 having a bipolar junction transistor 294 (e.g., a PNP-type bipolar junction transistor). The X-ray detector sub-module 292 includes a first surface 296 (e.g., a front surface) and a second surface 298 (e.g., a back surface) opposite the first surface 296. As shown, the bipolar junction transistor 272 includes an emitter 300 formed from a first P-type semiconductor layer 302 (e.g., P), a base 304 formed from an N-type semiconductor layer 306 (e.g., N-), and a collector 308 of a second P-type semiconductor layer 310 (e.g., P). The bipolar junction transistor 294 also includes a second N-type semiconductor layer 312 (e.g., N) disposed between the base 304 and the emitter 300. The second N-type semiconductor layer 312 is doped more heavily than the base 304 to avoid punch-through. The collector 308 is located on the first surface 296, and the emitter 300 is located on the second surface 298. The collector 308 is grounded. The bipolar junction transistor 294 functions as a detector element or active pixel. Electron-hole pairs (i.e., a current pulse) are generated in response to X-ray photons passing through the X-ray detector submodule 292 scattering valence electrons in a region, exciting the valence electrons into the conduction band. The amplified holes are injected into the base 304 (e.g., in an adjacent region on the second or back surface). The bipolar junction transistor 294 generates an output signal. The output signal is an amplified version of the current pulse generated by the active pixel in response to X-ray photons traveling through the semiconductor substrate.

[0058] Technical effects of the disclosed subject matter include providing a system and method for implementing an efficient active pixel sensor for a photon counting system to avoid placing a large capacitive load on the input CSA of the ASIC, which would increase system noise and power consumption. Technical effects also include improving the performance of silicon-based photon counting detectors (such as computed tomography detectors or other suitable types of radiography detectors). Technical effects further include improving image quality by reducing the effects of electronic noise (by improving the signal-to-noise ratio). Technical effects further include reducing system complexity and cost by lowering the power consumption of the readout circuitry (e.g., the ASIC).

[0059] The technology presented and claimed herein refers to and applies to specific examples and tangible objects that have practical properties that clearly improve the art, and as such is not abstract, intangible, or purely theoretical. Moreover, when a claim contains one or more elements designated as "means for [performing] ... a [function]" or "steps for [performing] ... a [function]," such elements are to be construed under 35 U.S.C. 112(f). However, for claims containing elements designated in other ways, such elements are not to be construed under 35 U.S.C. 112(f).

[0060] This description uses examples to disclose the present subject matter, including the best mode, and also enables any person skilled in the art to practice the subject matter, including making and using any devices or systems and performing any methods incorporating the same. The patentable scope of the present subject matter is defined by the claims, and may include other examples that occur to those skilled in the art. Such other examples are intended to be within the scope of the claims if they have structural elements that do not differ from the literal language of the claims, or if they include equivalent structural elements that do not differ in material way from the literal language of the claims. [Explanation of symbols]

[0061] 10 CT imaging systems 12 X-ray source 20 X-ray beams 22 Collimator 24 Subjects 26 X-ray 28 detector array 30 System Controller 32 Linear Positioning Subsystem 34 Rotation Subsystem 36 Motor Controller 38 X-ray controller 40 Data Collection System 42 Computer 44 Processing circuit 46 memory 48 Operator Workstations 50 displays 52 Printers 56 Remote Systems 58 detector submodules 59 Edge 60 detector elements 62 detector strips 64 segments 65 segments 66 Second Segment 68 Third Segment 69 Side edge 70 X-ray detector submodules 72 Semiconductor layer 78 Doped Implants 82 Gap 84 Electrical insulating layer 86 Wiring Traces 88 Passivation Layer 90 X-ray detector submodules 92 Diode 94 Voltage Bias 96 Trace 98 Readout circuit 102 Noise 104 detector capacity 108 X-ray detector submodules 110 X-ray detector submodule 112 Diode 114 Voltage Bias 116 Amplification stage 116 Photodiode 118 Trace 120 Readout circuit 124 Noise 125 detector capacity 126 X-ray detector submodule 128 Diode 130 Voltage Bias 132 Amplification stage 134 Transconductance Amplifier 136 transistors 138 Reset Circuit 140 pixel capacitor 142 Trace 144 Readout circuit 146 Current source 148 detector capacity 150 X-ray detector submodules 152 Diode 154 Voltage Bias 156 Amplification stage 158 Charge-sensitive amplifier 160 first transistor 161 Second Transistor 162 Current source 166 Feedback Capacitor 168 Trace 170 Readout circuit 172 detector capacity 174 X-ray detector submodule 176 Diode 178 Voltage Bias 180 Amplification stage 182 Transconductance Amplifier 184 transistors 186 Bias Reset Circuit 188 pixel capacitors 190 Trace 190 Readout circuit 192 Readout circuit 193 Current source 194 detector capacity 196 X-ray detector submodules 198 Diodes 200 Voltage Bias 202 Amplification stage 204 Charge Sensitive Amplifier 206 First Transistor 208 Second Transistor 210 Current source 214 Feedback Capacitor 216 Trace 218 Readout circuit 220 detector capacity 222 Method 224 blocks 226 blocks 228 blocks 232 Bipolar Junction Transistor 233 Voltage Bias 234 Trace 236 Readout circuit 240 Noise 242 detector capacity 244 X-ray detector submodule 246 Bipolar Junction Transistor 248 First Side 250 Second Side 252 Emitter 254 N-type semiconductor layer 256-based 258 P-type semiconductor layer 260 Collector 262 N-type semiconductor layer 264 N-type semiconductor layer 266 Junction 268 Junction 270 X-ray detector submodule 272 Bipolar Junction Transistor 274 First Side 276 Second Side 278 Emitter 280 N-type semiconductor layer 282 base 284 P-type semiconductor layer 286 Collector 288 N-type semiconductor layer 290 P-type semiconductor layer 292 X-ray detector submodule 294 Bipolar Junction Transistor 296 First Side 298 Second Side 300 emitters 302 P-type semiconductor layer 304 Base 306 N-type semiconductor layer 308 Collector 310 P-type semiconductor layer 312 N-type semiconductor layer

Claims

1. A photon-counting detector (28), a plurality of detector sub-modules (58), each detector sub-module (58) of the plurality of detector sub-modules (58) comprising: a semiconductor substrate (72); a plurality of active pixels (60) disposed on the semiconductor substrate (72) and configured to function as detector elements; and a plurality of traces (118) extending from the plurality of active pixels (60) to a readout circuit (120), each active pixel (60) of the plurality of active pixels (60) being coupled to a respective trace (118) of the plurality of traces (118), each active pixel (60) including an amplifier stage (116) configured to generate an output signal based on a current pulse output generated by the active pixel (60); a plurality of detector sub-modules, including: a readout circuit (120) configured to read out the output signals directly from the plurality of active pixels (60), the readout circuit (120) being separate from the plurality of detector sub-modules (58); a photon-counting detector,

2. 10. The photon-counting detector of claim 1, wherein the amplifier stage is configured to generate an output signal having a higher gain than the current pulse output to increase the signal-to-noise ratio.

3. The photon-counting detector (28) of claim 1, wherein each active pixel (60) is configured to function as a passive integrator.

4. 4. The photon-counting detector of claim 3, wherein the amplification stage of each active pixel includes a transconductance amplifier.

5. 5. The photon-counting detector of claim 4, wherein the amplification stage of each active pixel includes a single transistor.

6. 6. The photon-counting detector of claim 5, wherein each active pixel includes a pixel capacitor and a bias reset circuit, the bias reset circuit providing a reset path for the pixel capacitor and configured such that no bias is supplied to the readout circuit.

7. 6. The photon counting detector of claim 5, further comprising a current source disposed in the readout circuit, the current source biasing the single transistor.

8. 2. The photon-counting detector (28) of claim 1, wherein the amplification stage (116) of each active pixel (60) includes a charge-sensitive amplifier (158).

9. 9. The photon-counting detector of claim 8, wherein the amplification stage of each active pixel includes a first transistor and a second transistor, the first transistor being a charge-sensitive amplifier and the second transistor being configured to convert a voltage output into a current.

10. The photon-counting detector (28) of claim 9, wherein each active pixel (60) includes a current source (162).

11. The photon-counting detector (28) of claim 1, wherein the photon-counting detector (28) is configured for use in a computed tomography imaging system (10).

12. A computed tomography (CT) imaging system (10), comprising: A photon-counting detector (28), At least one detector sub-module (58), said at least one detector sub-module (58) comprising: a semiconductor substrate (72); a plurality of active pixels (60) disposed on the semiconductor substrate (72) and configured to function as detector elements; a plurality of traces (118) extending from the plurality of active pixels (60) to a readout circuit (120), each active pixel (60) of the plurality of active pixels (60) being coupled to a respective trace (118) of the plurality of traces (118), each active pixel (60) including an amplifier stage (116) configured to generate an output signal based on a current pulse output generated by the active pixel (60); at least one detector sub-module including: a readout circuit (120) configured to directly read out the output signals from the plurality of active pixels (60), the readout circuit (120) being separate from the at least one detector sub-module (58). a photon counting detector including A CT imaging system comprising:

13. The CT imaging system (10) of claim 12, wherein each active pixel (60) is configured to function as a passive integrator.

14. 14. The CT imaging system (10) of claim 13, wherein the amplification stage (116) of each active pixel (60) includes a transconductance amplifier (134).

15. 15. The CT imaging system of claim 14, wherein the amplifier stage (116) of each active pixel (60) includes a single transistor (136).

Images