JP2020038185A - Photon counting detector and X-ray CT apparatus - Google Patents

Abstract

The present invention is capable of performing well in both high flux and low flux situations. A photon counting detector includes a plurality of macropixels arranged in a semiconductor crystal having a first surface and a second surface parallel to the first surface. Each macropixel of the plurality of macropixels collects projection data for generating a reconstructed image. Each macro-pixel is of only two types, at least one large micro-pixel arranged within each macro-pixel and at least two small micro-pixels arranged within each macro-pixel. The area of each of the at least two small micropixels is smaller than the area of the at least one large micropixel. [Selection drawing] Fig. 4B

Description

  Embodiments of the present invention relate to a photon counting detector and an X-ray CT apparatus.

  Photon counting computed tomography (CT) is a computed tomography technique that has the potential to significantly improve existing CT imaging techniques. Photon counting CT systems include photon counting detectors that record the interaction of individual photons. By tracing the energy stored in each interaction, the detector pixels of the photon-counting CT detector each record an approximate energy spectrum or count in an energy bin. In contrast, a typical CT scanner uses an energy integrating detector in which the total energy stored in a pixel over a fixed period of time is recorded.

There are several problems associated with incorporating photon counting detectors into CT systems. Such problems relate to the requirements for detector materials and electronics resulting from large data volumes and count rates. As an example, each mm ² of the CT detector can receive millions of photon interactions every second during a scan. The pulse resolution time should be small compared to the average time between photon interactions at the pixel to avoid saturation in the area where there is small material between the X-ray source and the detector It is. Even before saturation, the functionality of the detector is due to pulse pile-up due to two (or more) photon interactions occurring at the same pixel being too close in time to separate as separate events. And it is starting to decline. Such quasi-coincided interactions lead to photon count loss and pulse waveform distortion. These effects greatly increase the demands on the physical response time of the detector material, as well as the electronics involved for pulse shaping, binning and recording of pixel data.

Photon counting detectors can be manufactured with pixel sizes even smaller than about 1 × 1 mm ² without compromising dose efficiency. Using smaller pixels reduces the count rate per pixel, thus exacerbating the requirement for pulse resolution time at the expense of requiring more electronics. Pixel size and corresponding readout electronics design are key parts in photon counting detector design. For high flux scanning environments that involve millions of photon interactions per second during a CT scan, smaller pixel sizes are preferred for photon counting detectors (PCDs). Generally, pixel sizes for photon counting CT applications can range from 150 μm to 500 μm. That is, if the pixel size is smaller than the pixel size, the pixel interval (pixel pitch) is closer. Smaller pixel designs may consist of fewer pixels compared to 500 μm size for photon counting CT applications. Smaller pixel designs for PCDs are desirable for high flux because they are relatively insensitive to pulse pile-up. However, one problem associated with the use of smaller pixels is severe charge sharing and poor detector response due to signal crosstalk.

  Partial energy storage and a single photon producing a signal at a large number of pixels pose challenges in photon counting CT. Charge sharing is one of such events in which interactions occur near pixel boundaries, and the emitted energy is shared between neighboring pixels, so that some energy is interpreted as lower photons. Is a factor. Charge sharing results in a distorted energy spectrum. In contrast to saturation and pile-up effects, the problem is that partial energy storage and large numbers of interacting photons are exacerbated by smaller pixel sizes.

  Using a larger pixel size for the PCD is one solution that provides less charge sharing and crosstalk effects based on the typically good detector response. The peak energy fraction that leaks into the lower energy fraction is smaller, which can lead to improvements to material discrimination. In this way, using large pixel sizes, such as 500 μm sized pixels, may be desirable in low flux scanning environments. One potential drawback of using a large pixel design for the PCD places a significant burden on electronics or other post-processing issues to address severe pulse pile-up. Thus, in terms of material discrimination noise, a small pixel size may be desirable in a high flux scan environment and a large pixel size may be desirable in a low flux scan environment. In a photon counting CT scan environment, the PCD is typically exposed to high and low flux during a patient scan. PCD in photon-counting CT scanners using small pixel designs may work well for high flux, but may perform poorly in low flux situations. Conversely, PCDs in photon counting CT scanners that use large pixel designs work well for low flux situations, but may perform poorly in high flux situations.

  Therefore, by minimizing the pulse pile-up problem associated with the use of large pixels in the PCD, as well as the charge sharing problem associated with small pixels in the PCD, both high flux and low flux CT scan environments However, there is a technical need for a fully functioning photon counting CT detector.

Japanese Patent No. 6355747

  The problem addressed by the present invention is that scans that can include both high and low fluxes that are relatively insensitive to pulse pile-up while maintaining sufficient detector response by reducing charge sharing It is directed to CT detector devices for use in the environment. CT detector devices can perform well in both high and low flux situations by using a photon counting detector (PCD) that implements a hybrid pixel pattern design.

  The photon counting detector according to the present embodiment includes a plurality of macropixels arranged on a semiconductor crystal having a first surface and a second surface parallel to the first surface. Each macro pixel of the plurality of macro pixels collects projection data for generating a reconstructed image. Each macro pixel is composed of only two types: at least one large micro pixel arranged in each macro pixel and at least two small micro pixels arranged in each macro pixel. The area of each of the at least two small micropixels is smaller than the area of the at least one large micropixel.

FIG. 1 shows a schematic diagram of an implementation example of a computed tomography (CT) scanner. FIG. 2 shows a schematic diagram of a uniform pixel pattern used in a photon counting detector. FIG. 4 shows a schematic diagram of another uniform pixel pattern used in a photon counting detector. 5 is a graph illustrating an example of two different detector responses for two different pixel sizes. 4 is a graph illustrating an example of a relationship between discrimination noise and pixel size under a high flux and a low flux. 1 is a perspective view of a photon counting detector including a plurality of pixels according to the present disclosure. FIG. 4B is a schematic diagram of a hybrid pixel pattern for each pixel in the photon counting detector of FIG. 4A according to the present disclosure. FIG. 4 is a schematic diagram of another hybrid pixel pattern for use in a photon counting detector according to the present disclosure. FIG. 5B is a schematic diagram of a group of six pixels incorporating the hybrid pixel pattern of FIG. 5A according to the present disclosure. FIG. 2 is a schematic diagram of a hybrid pixel pattern for use in a photon counting detector according to the present disclosure. 1 is a diagram illustrating a configuration example of an X-ray detector according to the present disclosure. FIG. 3 shows a schematic diagram of a hybrid pixel pattern design and corresponding input channels from the pixels of the amplifier and counter stages to reduce charge sharing according to the present disclosure. FIG. 3 shows a schematic diagram of a hybrid pixel pattern design and corresponding pixels of the amplifier and counter stages to minimize the pulse pile-up problem according to the present disclosure. FIG. 4 shows a schematic diagram of a hybrid pixel pattern design and corresponding amplifier and counter stage pixel to input channels according to the present disclosure.

  The following description is of specific embodiments, but other embodiments may include alternatives, equivalents, and variations. In addition, embodiments may include some novel features, and particular features may not be essential for performing the devices, systems, and methods described herein.

  The present disclosure focuses on a photon counting detector (PCD) used in a photon counting detector CT system. The term "PCD" is used synonymously with the term "detector" throughout this disclosure. When the photon counting CT system scans a patient or subject, the scanning environment may change continuously from high flux to low flux, or vice versa. PCDs with the hybrid pixel pattern design discussed throughout this disclosure provide charge sharing even during dynamic changes in the scan environment, where the PCD is exposed to both high and low flux scan environments. By reducing this, it is relatively insensitive to pulse pile-up while maintaining sufficient detector response.

The uniform pixel pattern design shown in FIG. 2A shows one macropixel 50 with 16 equally sized micropixels 52. If the area of the macropixel 50 is 1 mm ² , the pixel size of the 16 equally sized micropixels is 250 μm. The detector response in this example has a pixel size of 250 μm, with the benefit of being less susceptible to pulse pile-up in high flux scan environments and the disadvantage of severe charge sharing in low flux scan environments. May exhibit related features. Also in FIG. 2B, one macropixel 60 includes a uniform pixel pattern design with four equally sized micropixels 62. If the area of the macro pixel 60 is the same as the macro pixel 50 in FIG. 2A, the pixel size of the four equally sized micro pixels is 500 μm. The detector response in this example exhibits features associated with a pixel size of 500 μm, such as the benefit of minimal charge sharing in a low flux scan environment and the disadvantage of increased pulse pile-up in a high flux scan environment. there is a possibility. Therefore, using only a uniform pixel pattern design corresponding to multiple macropixels on the PCD may not be well suited for a scanning environment that includes both high flux and low flux.

  The hybrid pixel pattern for each macropixel in the PCD described throughout this disclosure provides a high and low flux scan environment by implementing a non-uniform pixel pattern design consisting of at least two differently sized micropixels. Designed. A hybrid pixel pattern design of macropixels can be implemented for the PCD included in a photon counting CT scan system as described above with reference to FIG.

  Hereinafter, an X-ray CT (Computed Tomography) apparatus including the medical image processing apparatus according to the present embodiment will be described with reference to the block diagram of FIG. The X-ray CT apparatus 1 illustrated in FIG. 1 includes a gantry device 10, a bed device 30, and a console device 40 that implements processing of a medical image processing device. In FIG. 1, for convenience of explanation, a plurality of gantry devices 10 are drawn.

  In this embodiment, the longitudinal direction of the rotation axis of the rotating frame 13 or the top plate 33 of the bed device 30 in the non-tilted state is perpendicular to the Z-axis direction and the Z-axis direction, and is an axial direction that is horizontal to the floor surface. Is defined as an axis direction orthogonal to the X-axis direction and the Z-axis direction and perpendicular to the floor surface as a Y-axis direction.

  For example, the gantry device 10 and the bed device 30 are installed in a CT examination room, and the console device 40 is installed in a control room adjacent to the CT examination room. Note that the console device 40 does not necessarily have to be installed in the control room. For example, the console device 40 may be installed in the same room with the gantry device 10 and the bed device 30. In any case, the gantry device 10, the bed device 30, and the console device 40 are communicably connected by wire or wirelessly.

  The gantry device 10 is a scanning device having a configuration for performing X-ray CT imaging of the subject P. The gantry device 10 includes an X-ray tube 11, an X-ray detector 12, a rotating frame 13, an X-ray high-voltage device 14, a control device 15, a wedge 16, a collimator 17, and a data collection device 18 (hereinafter, referred to as a data collection device). , DAS (Data Acquisition System) 18).

  The X-ray tube 11 is a vacuum tube that generates X-rays by irradiating thermoelectrons from a cathode (filament) to an anode (target) by applying a high voltage from an X-ray high voltage device 14 and supplying a filament current. It is. Specifically, X-rays are generated by the collision of the thermal electrons with the target. For example, the X-ray tube 11 includes a rotating anode type X-ray tube that generates X-rays by irradiating a rotating anode with thermoelectrons. The X-rays generated by the X-ray tube 11 are formed into a cone beam shape via, for example, a collimator 17 and are irradiated on the subject P.

  The X-ray detector 12 detects X-rays emitted from the X-ray tube 11 and passed through the subject P, and outputs an electric signal corresponding to the X-ray dose to the DAS 18. The X-ray detector 12 has, for example, a plurality of X-ray detection element rows in which a plurality of X-ray detection elements are arranged in the channel direction along one arc around the focal point of the X-ray tube 11. The X-ray detector 12 has, for example, a row structure in which a plurality of X-ray detection element rows in which a plurality of X-ray detection elements are arrayed in a channel direction are arranged in a slice direction (row direction, row direction).

  Specifically, the X-ray detector 12 may be, for example, an indirect conversion type detector having a grid, a scintillator array, and an optical sensor array, or an incident X-ray detector as described later with reference to FIG. May be a direct conversion type detector having a semiconductor element for converting the signal into an electric signal. The X-ray detector 12 is an example of the PCD according to the present embodiment, and is hereinafter also referred to as the PCD 12.

  The scintillator array has a plurality of scintillators. The scintillator converts incident X-rays into photons of a number corresponding to the intensity of the incident X-rays.

  The grid has an X-ray shielding plate disposed on the surface of the scintillator array on the X-ray incident side and having a function of absorbing scattered X-rays. The grid may be called a collimator.

  The optical sensor array has a function of amplifying light received from the scintillator, converting the light into an electric signal, and generating an output signal (energy signal) having a peak value according to the energy of the incident X-ray. It has an optical sensor such as a photomultiplier tube (photomultiplier: PMT).

  The rotating frame 13 supports the X-ray generator and the X-ray detector 12 so as to be rotatable around a rotation axis. Specifically, the rotating frame 13 supports the X-ray tube 11 and the X-ray detector 12 so as to face each other, and rotates the X-ray tube 11 and the X-ray detector 12 by a control device 15 described later. It is. The rotating frame 13 is rotatably supported by a fixed frame (not shown) made of metal such as aluminum. Specifically, the rotating frame 13 is connected to an edge of the fixed frame via a bearing. The rotating frame 13 receives power from the driving mechanism of the control device 15 and rotates around the rotation axis Z at a constant angular velocity.

  The rotating frame 13 further includes and supports an X-ray high-voltage device 14 and a DAS 18 in addition to the X-ray tube 11 and the X-ray detector 12. Such a rotating frame 13 is housed in a substantially cylindrical housing in which an opening (bore) 19 forming an imaging space is formed. The opening substantially matches the FOV. The central axis of the opening coincides with the rotation axis Z of the rotation frame 13. The detection data generated by the DAS 18 is provided on a non-rotating portion (for example, a fixed frame, not shown in FIG. 1) of the gantry device by optical communication from a transmitter having, for example, a light emitting diode (LED). The data is transmitted to a receiver (not shown) having a photodiode, and is transmitted to the console device 40. Note that the method of transmitting the detection data from the rotating frame to the non-rotating portion of the gantry is not limited to the above-described optical communication, and any method may be adopted as long as it is a non-contact type data transmission.

  The X-ray high-voltage device 14 has an electric circuit such as a transformer (transformer) and a rectifier, and has a function of generating a high voltage applied to the X-ray tube 11 and a function of generating a filament current supplied to the X-ray tube 11. It has a generator and an X-ray controller for controlling the output voltage according to the X-ray emitted by the X-ray tube 11. The high voltage generator may be of a transformer type or an inverter type. The X-ray high-voltage device 14 may be provided on a rotating frame 13 described later, or may be provided on a fixed frame (not shown) of the gantry device 10.

  The control device 15 has a processing circuit having a CPU (Central Processing Unit) and the like, and a driving mechanism such as a motor and an actuator. The processing circuit has, as hardware resources, a processor such as a CPU and an MPU (Micro Processing Unit) and a memory such as a ROM (Read Only Memory) and a RAM (Random Access Memory). In addition, the control device 15 includes an application specific integrated circuit (ASIC), a field programmable gate array (FPGA), and another complex programmable logic device (CPLD). ), A simple programmable logic device (SPLD). The control device 15 controls the X-ray high-voltage device 14, the DAS 18, and the like according to a command from the console device 40. The processor realizes the above control by reading and implementing a program stored in the memory.

  A CPU may execute a computer program that includes a set of computer-readable instructions that perform the functions described herein, such a computer program comprising any of the non-transitory electronic memory and / or hard disk drives described above. , CD (Compact Disc), DVD (Digital Versatile Disc), FLASH drive, or any other known storage medium. Further, the computer readable instructions may be provided as a utility application, a background daemon, or a component of an operating system, or a combination thereof, with other operating systems known to those skilled in the art executing together with the processor. . Further, the CPU may be implemented as a multiple processor, operating in parallel and cooperatively to execute instructions.

  In addition, the control device 15 has a function of receiving an input signal from an input interface 43, which will be described later, attached to the console device 40 or the gantry device 10, and controlling the operation of the gantry device 10 and the couch device 30. For example, the control device 15 performs control to rotate the rotating frame 13 in response to the input signal, control to tilt the gantry device 10, and control to operate the bed device 30 and the top board 33. The control for tilting the gantry device 10 is performed by the control device 15 based on the tilt angle (tilt angle) information input by the input interface 43 attached to the gantry device 10, with the rotation frame around the axis parallel to the X-axis direction. 13 is rotated. Further, the control device 15 may be provided in the gantry device 10 or may be provided in the console device 40. Note that the control device 15 may be configured to directly incorporate the program into the circuit of the processor instead of storing the program in the memory. In this case, the processor realizes the above control by reading and executing a program incorporated in the circuit.

  The wedge 16 is a filter for adjusting the X-ray dose emitted from the X-ray tube 11. More specifically, the wedge 16 transmits and attenuates the X-rays emitted from the X-ray tube 11 so that the X-rays emitted from the X-ray tube 11 to the subject P have a predetermined distribution. Filter. For example, the wedge 16 (wedge filter, bow-tie filter) is a filter formed by processing aluminum so as to have a predetermined target angle and a predetermined thickness.

  The collimator 17 is a lead plate or the like for narrowing an irradiation range of the X-ray transmitted through the wedge 16, and forms a slit by a combination of a plurality of lead plates and the like. Note that the collimator 17 is sometimes called an X-ray aperture.

  The DAS 18 generates digital data indicating the count of X-rays detected by the X-ray detector 12 (hereinafter, also referred to as detection data) for each of a plurality of energy bands (hereinafter, also referred to as energy bins or simply bins). I do. The detection data is a set of data including a channel number, a column number, a view number indicating a collected view (also referred to as a projection angle) of the source X-ray detection element, and a count value identified by an energy bin number. . The DAS 18 is realized by, for example, an ASIC (Application Specific Integrated Circuit) equipped with a circuit element capable of generating detection data. The detection data is transferred to the console device 40.

  For example, DAS 18 includes a preamplifier, a variable amplifier, an integrator, and an A / D converter for each detector pixel. The preamplifier amplifies an electric signal from the connection source X-ray detection element with a predetermined gain. The variable amplifier amplifies the electric signal from the preamplifier with a variable gain. The integrating circuit integrates the electric signal from the preamplifier over one view period to generate an integrated signal. The peak value of the integration signal corresponds to the X-ray dose value detected by the connection source X-ray detection element over one view period. The A / D converter performs analog-to-digital conversion of the integration signal from the integration circuit to generate detection data.

  The couch device 30 is a device for placing and moving the subject P to be scanned, and includes a base 31, a couch driving device 32, a top plate 33, and a support frame.

  The base 31 is a housing that supports the support frame 34 movably in the vertical direction.

  The couch driving device 32 is a motor or an actuator that moves the table 33 on which the subject P is placed in the longitudinal direction of the table 33. The couch driving device 32 moves the couchtop 33 according to control by the console device 40 or control by the control device 15. For example, the couch driving device 32 moves the table 33 in a direction perpendicular to the object P such that the body axis of the object P placed on the table 33 coincides with the central axis of the opening of the rotating frame 13. I do. Further, the couch driving device 32 may move the couchtop 33 along the body axis direction of the subject P in accordance with X-ray CT imaging performed using the gantry device 10. The couch driving device 32 generates power by being driven at a rotation speed corresponding to a duty ratio of a driving signal from the control device 15 or the like. The couch driving device 32 is realized by a motor such as a direct drive motor or a servo motor.

  The top plate 33 provided on the upper surface of the support frame 34 is a plate on which the subject P is placed. The bed driving device 32 may move the support frame 34 in the longitudinal direction of the top plate 33 in addition to the top plate 33.

  The console device 40 has a memory 41, a display 42, an input interface 43, and a processing circuit 44. Data communication among the memory 41, the display 42, the input interface 43, and the processing circuit 44 is performed via a bus (BUS). Although the console device 40 is described as being separate from the gantry device 10, the gantry device 10 may include the console device 40 or some of the components of the console device 40.

  The memory 41 is a storage device such as a hard disk drive (HDD), a solid state drive (SSD), or an integrated circuit storage device for storing various information. The memory 41 stores, for example, projection data and reconstructed image data. The memory 41 is a drive device that reads and writes various information between a portable storage medium such as a CD, a DVD, and a flash memory, and a semiconductor memory element such as a RAM (Random Access Memory) in addition to the HDD and the SSD. It may be. Further, the storage area of the memory 41 may be in the X-ray CT apparatus 1 or in an external storage device connected via a network. For example, the memory 41 stores data of a CT image and a display image. Further, the memory 41 stores a control program according to the present embodiment.

  The display 42 displays various information. For example, the display 42 outputs a medical image (CT image) generated by the processing circuit 44, a GUI (Graphical User Interface) for receiving various operations from the operator, and the like. For example, as the display 42, for example, a liquid crystal display (LCD: Liquid Crystal Display), a CRT (Cathode Ray Tube) display, an organic EL display (OELD: Organic Electro Luminescence Display), a plasma display, or any other display is appropriately used. , Has become available. The display 42 may be provided on the gantry device 10. Further, the display 42 may be a desktop type, or may be configured by a tablet terminal or the like that can wirelessly communicate with the console device 40 main body.

  The input interface 43 receives various input operations from the operator, converts the received input operations into electric signals, and outputs the electric signals to the processing circuit 44. For example, the input interface 43 receives, from the operator, acquisition conditions for acquiring projection data, reconstruction conditions for reconstructing a CT image, image processing conditions for generating a post-processed image from a CT image, and the like. . As the input interface 43, for example, a mouse, a keyboard, a trackball, a switch, a button, a joystick, a touch pad, a touch panel display, and the like can be appropriately used. Note that, in the present embodiment, the input interface 43 is not limited to one having physical operation components such as a mouse, a keyboard, a trackball, a switch, a button, a joystick, a touch pad, and a touch panel display. For example, an example of the input interface 43 includes an electric signal processing circuit that receives an electric signal corresponding to an input operation from an external input device provided separately from the device and outputs the electric signal to the processing circuit 44. . The input interface 43 may be provided in the gantry device 10. Further, the input interface 43 may be configured by a tablet terminal or the like that can wirelessly communicate with the console device body.

  The processing circuit 44 controls the overall operation of the X-ray CT apparatus 1 in accordance with an input operation electric signal output from the input interface 43. For example, the processing circuit 44 has, as hardware resources, a processor such as a CPU, an MPU, and a GPU (Graphics Processing Unit) and a memory such as a ROM and a RAM. The processing circuit 44 executes a system control function 441, a pre-processing function 442, a reconstruction processing function 443, and a display control function 444 by a processor that executes a program expanded in the memory. Note that each function (the system control function 441, the preprocessing function 442, the reconfiguration processing function 443, and the display control function 444) is not limited to being realized by a single processing circuit. A processing circuit may be configured by combining a plurality of independent processors, and each processor may execute a program to realize each function.

  The system control function 441 controls each function of the processing circuit 44 based on an input operation received from the operator via the input interface 43. Specifically, the system control function 441 reads out the control program stored in the memory 41, expands it on the memory in the processing circuit 44, and controls each unit of the X-ray CT apparatus 1 according to the expanded control program. . For example, the processing circuit 44 controls each function of the processing circuit 44 based on an input operation received from the operator via the input interface 43. For example, the system control function 441 acquires a two-dimensional positioning image of the subject P for determining a scan range, imaging conditions, and the like. Note that the positioning image is also called a scano image or a scout image.

  The pre-processing function 442 performs logarithmic conversion processing, offset correction processing, sensitivity correction processing between channels, beam hardening correction, detector calibration, detector non-linearity, polarity effect, noise on the detection data output from the DAS 18. Generates data that has been subjected to preprocessing such as balancing and correction for material discrimination. The data before the pre-processing (detection data) and the data after the pre-processing may be collectively referred to as projection data. The pre-processing function 442 is an example of a pre-processing unit.

  The reconstruction processing function 443 performs a reconstruction process on the projection data generated by the preprocessing function 442 using a filter-corrected backprojection method, an iterative approximation reconstruction method, a stochastic image reconstruction method, or the like. To generate CT image data. The reconfiguration processing function 443 is an example of a reconfiguration processing unit. The CT image data may be subjected to image filtering, smoothing, volume rendering processing, and image difference processing as needed.

The display control function 444 converts the CT image data generated by the reconstruction processing function 443 based on an input operation received from the operator via the input interface 43 into layer image data or three-dimensional data of an arbitrary cross section by a known method. Convert to image data. Note that the generation of the three-dimensional image data may be directly performed by the reconstruction processing function 443. The display control function 444 is an example of a display control unit.
In the present embodiment, the “projection plane” indicates a volume through which X-rays pass from the X-ray tube 11 to the PCD 12. The “subject space” refers to a place where the projection plane intersects with the opening of the gantry. “Image space” includes the integration of projection planes corresponding to all projection angles of X-ray tube 11 as X-ray tube 11 rotates around the opening of the gantry. The image space is usually larger than the object space, since the image space allows image reconstruction for a volume that is wider than the size of the object P.

  When the subject P occupies the subject space and the gantry 10 rotates through a series of projection angles while the gantry 10 collects X-ray transmission / attenuation projection data through the subject P at each projection angle. , A scan is performed.

Generally, PCD 12 outputs the number of photons for each of a predetermined number of energy bins. In one implementation, the PCD 12 is sparsely positioned around the subject P in a predetermined geometry.
In one implementation, X-ray tube 11 is a single source that emits a broad spectrum of X-ray energy.

PCD12 is a direct conversion type X-ray radiation detection based on semiconductors such as cadmium telluride (CdTe), cadmium zinc telluride (CZT), silicon (Si), mercury iodide (HgI ₂ ), and gallium arsenide (GaAs). You can use a container. A direct conversion type X-ray detector based on a semiconductor has a much faster time response than an indirect conversion type detector such as a scintillator detector. This fast time response of the direct conversion detector allows the detector to resolve individual X-ray detector events. However, at high x-ray flux typical of clinical x-ray applications, pile-up can occur in some of the detection events. The energy of the detected X-rays is proportional to the signal generated by the direct conversion detector, and the detection events are organized into energy bins that produce spectrally resolved X-ray data for the spectral CT. be able to.

  In the present disclosure, the signal from the PCD 12 is controlled by a read circuit (not shown). The read channel outputs an analog signal processed by the amplifier, which is then input to an analog-to-digital (A / D) converter. Such an A / D converter is clocked at a high frequency (generally 40 to 100 MHz) and outputs a digital signal. Such digital signals are then transferred to a processor for detection of incident photons. The digital samples are processed by an algorithm implemented in a digital filter that determines the time stamp and amplitude of the signal. The readout circuit continually samples the analog signal and produces a large data set for further processing in the image processing chain.

  The disclosed embodiments facilitate the use of thinner detectors (eg, W <2.0 mm) at larger thicknesses (T-3.0 mm). Such a detector improves charge collection efficiency while maintaining a higher operating area without charge loss at the surface. The weighted potential describes the coupling between the moving charge and the electrodes of the charge as the moving charge drifts at the detector.

FIG. 3A is a graph illustrating an example x-ray response at 100 kiloelectron volts (keV) for a PCD that includes a uniform pixel pattern design using pixel sizes at 500 μm and 250 μm pixels, respectively. Solid lines represent PCDs with 250 μm pixels, and dashed lines represent PCDs with 500 μm pixels. Both PCDs are, for example, 2 mm thick and have macro pixels with an area of 1 mm ² . A macropixel having a uniform pixel pattern design with a pixel size of 500 μm may include four micropixels in a macropixel with an area of 1 mm ² . Also, a macropixel having a uniform pixel pattern design with a pixel size of 250 μm may include 16 micropixels in a macropixel having an area of 1 mm ² . The dashed line represents a PCD with 500 μm uniform pixels. Two different pixel sizes, 250 μm and 500 μm, are shown to indicate different detector responses due to pixel size. The vertical axis of the graph shown in FIG. 3A reflects the count or percentage detected at a particular energy for a 100 keV input. The graph shows that the 250 μm pixel has a higher charge sharing event. An ideal detector would have a detector response that represented a delta function. The graph shows how pixel size and corresponding readout electronics are important for PCD design.

  A PCD with a uniform pixel pattern design with smaller micropixels in the example of FIG. 3A is desirable for high flux scanning environments, which may consist of millions of photon interactions per square millimeter per second. One advantage associated with smaller micropixels is that they are relatively insensitive to pulse pile-up problems. However, due to the possible severe charge sharing as shown in FIG. 3, smaller pixel sizes appear to be more susceptible to detector response degradation. Pixel sizes as large as 500 μm in this example can provide lower charge sharing and crosstalk effects, and can result in improved detector response. Part of the peak energy leakage to the lower energy part is smaller, which makes it possible to reduce the material discrimination noise. However, larger micropixel designs increase the burden on electronics or other post-processing to deal with severe pulse pile-up.

  In terms of material discrimination noise, as shown in the graph illustrated in FIG. 3B, small pixel sizes are preferred in a high flux scan environment, and large pixel sizes are preferred in a low flux scan environment. The graph in FIG. 3B illustrates the relationship between discrimination noise and pixel size in a high flux and low flux scan environment. The vertical axis of the graph in FIG. 3B represents discrimination noise, and the horizontal axis represents a pixel size range of 250-500 μm. The black circles reflect the level of discrimination noise in a low flux scan environment as the pixel size increases, and the white circles reflect the level of discrimination noise in a high flux scan environment as the pixel size increases. According to the graph of FIG. 3B, each circle reveals that the discrimination noise also increases in a high flux environment where the pixel size increases, and that the discrimination noise decreases as the pixel size increases in the low flux scan. ing. In this way, using a detector with a uniform pixel size design, including small or large pixels, may provide certain advantages and disadvantages.

  Referring now to FIG. 4A, a photon counting detector array 70 including a plurality of macropixels in a hybrid pixel pattern design is shown. The photon counting detector array 70 includes a crystal 71 (semiconductor crystal) formed of a semiconductor material such as CdZnTe or CdTe. One face of the crystal 71 has a large single cathode electrode 76. The opposite side of the crystal is the anode surface 72, which includes a rectangular or square anode pixel 74. Hybrid pixel pattern design to enjoy the benefits of using small pixel size designs as well as large pixel size designs while minimizing any disadvantages associated with using either small pixel size designs or large pixel size designs Is used. The thickness of the detector array 70 is denoted by T. In one example of the present disclosure, the thickness of the detector array 70 is 3 mm, which is sufficient to stop X-ray power. It should be noted that the thickness of the detector array 70 may vary depending on the type of application, semiconductor material, and other considerations associated with the detector design. Each individual pixel 74 of the detector array 70 is a macro pixel in a hybrid pixel pattern design. Each macropixel in the detector array 70 is of uniform size, while macropixels include micropixels of various sizes. For example, the detector may be a 12 × 8 array of macropixels (96 macropixels), a 12 × 12 array of macropixels (144 macropixels), or a 16 × 16 array of macropixels (256 macropixels). , But is not intended to limit the array of macropixels associated with a particular detector.

  In one example, the macropixels of detector array 70 shown in FIG. 4A have an area of 1 mm × 1 mm. The area of the macropixel can be of any size and is not restricted to an area of 1 mm × 1 mm.

The present disclosure minimizes pulse pile-up and charge sharing problems by using a hybrid pixel pattern design consisting of non-uniform micro-pixels in macro-pixels as shown by way of example in FIG. 4B. FIG. 4B includes a macro pixel 74 with a hybrid pixel pattern design. Macro pixel 74 includes two different sized micro pixels. In this example, one large micropixel 82 (also called a large micropixel) in the macropixel 74 is surrounded by twelve smaller micropixels 80 (also called small micropixels). The macro pixel 74 in this example is composed of a total of 12 micro pixels. If the macropixel 74 in this example has an area of 1 mm × 1 mm, then the large micropixel 82 has an area four times larger than the area of one of the small micropixels 80. The smaller micropixel in this example has a 250 μm size and the larger micropixel has a 500 μm size. The hybrid pixel pattern design for the macropixel 74 substantially relies on a single larger 500 μm micropixel 82 in a low flux scanning environment to minimize charge sharing and crosstalk between the micropixels of the macropixel 74. , May rely substantially on 250 μm micropixels 80 over a high flux scanning environment to avoid pulse pile-up problems. Large micropixels 82 are expected to provide better information under a low flux scan environment, and small micropixels 80 are each expected to provide better information under a high flux scan environment. .
In other words, information obtained from the first set of detector responses for at least one large micropixel 82 may be used to calibrate at least two small micropixels 80 in a low flux scan environment. Conversely, information obtained from the second set of detector responses for at least two small micropixels 80 may be used to calibrate at least one large micropixel 80 in a high flux scanning environment. The detector response for both large and small micropixels in a macropixel is used regardless of the type of scanning environment (high flux / low flux).

  Unlike a uniform pixel pattern design for macropixels, a hybrid pixel pattern design can maintain a good balance between large and small micropixels. From the Shockley-Ramo theorem, the weighting region and weighting potential for each macropixel will be determined by the micropixel size itself. A hybrid pixel pattern is a detector pattern that includes pixels of two or more sizes for the purpose of optimizing the detector's spectral photon counting performance over a range of input flux rates. In the above example, the hybrid pixel pattern for a macro pixel includes two different sized micro pixels. However, the macropixels of the hybrid pixel pattern design may include three different sized micropixels, or four different sized micropixels. The detector response increases in combination with three or four differently sized micropixels in the macropixel.

  Although small micropixels have been described as being 250 μm in size, various sizes, such as 225 μm, 200 μm, 175 μm, 150 μm, etc., can be used for smaller micropixels, from about 150 μm to about 300 μm. It may be a size between. Typically, sizes of 250 μm or less can be used for smaller micropixel size designs associated with the desired results in a high flux scanning environment.

  Although the large micropixel is described as having a size of 500 μm, various sizes, for example, 450 μm, 400 μm, 350 μm, and 300 μm can be used for a large micropixel. Sizes up to 600 μm may be used.

  The size is set so that the large micropixel and the small micropixel have a similar relationship. For example, for a large micropixel and a small micropixel, half the length of one side of the macropixel is equal to at least one side of the large micropixel, and half the length of one side of the large micropixel is the small micropixel. May have a relationship corresponding to the length of at least one side. Thus, a large micropixel and a small micropixel may have a self-similar relationship. Large and small micropixels having a self-similar relationship are flat-filled on the detection surface of the X-ray detector 12.

  In general, there is no clear rule on the minimum pixel size for small micropixels. However, the charge sharing effect increases as the pixel size decreases. Pixel sizes smaller than the 150 μm size may not be of interest for CT applications. Each micropixel has an electronics readout channel to handle signals from a particular micropixel. Application specific integrated circuits (ASICs) are often used for such applications. Each electronics channel is typically a pre-amplifier, a molding machine (shaper, CR-RC shaper, etc.), or a high / low-pass filter to record the event amplitude (analog-to-digital converted or implemented by a comparator). And a digitizer. For most read electronics, eg, for ASICs, all read channels are typically the same physical size. Readout channels of different physical sizes may be used to match the hybrid pixel pattern design of the detector, or readout channels of the same physical size may be used with the help of an interposer Is also good.

  Referring to FIG. 5A, a macro pixel 90 having a hybrid pixel pattern design different from the design shown in FIG. 4B is illustrated. The macro pixel 90 includes two large micro pixels 92 and eight small micro pixels 94. In this example, a macro pixel 90 having an area of 1 mm × 1 mm has two large micro pixels with a pixel size of 500 μm and eight smaller micro pixels with a pixel size of 250 μm. The macro pixels 90 include a total of 10 micro pixels in the design. The hybrid pixel design still enjoys the benefits of using two different sized micropixels in the macropixel 90 for both high flux and low flux scanning environments. In addition, the use of two large micropixels and eight smaller micropixels results in an image with better resolution, thanks to the addition of larger micropixels.

  FIG. 5B shows a group of six macropixels 100, including the hybrid pixel pattern design of macropixels 90 shown in FIG. 5A. A group of six macropixels is an example of which parts of the photon counting detector are similar on the anode side.

  Referring to FIG. 5C, a macropixel 91 having a hybrid pixel pattern design different from the designs shown in FIGS. 5A and 5B is shown. Such a macropixel 91 includes three large micropixels 93 and four smaller micropixels 95. In this example, in a macropixel 91 having an area of 1 mm × 1 mm, three large micropixels have a pixel size of 500 μm and four small micropixels have a pixel size of 250 μm. The macro pixel 91 includes a total of seven micro pixels in the design. Hybrid pixel pattern designs also benefit from using different sized micropixels in macropixels 91 for both high dose and low flux scan environments. The use of three large micropixels and four small micropixels results in an image with better resolution.

The high flux scan environment includes a fast energy resolved photon-counting x-ray imaging array using a cadmium telluride (CdTe) semiconductor sensor composed of pixels and measured with a clinical CT x-ray source Output count rate (OCR) of over 100 million counts per second per square millimeter (Mcps / mm ² ). High-speed application-specific integrated circuits (ASICs) are two-dimensional (two-dimensional) arrays for reading from CdTe sensors composed of pixels and cadmium zinc telluride (CdZnTe) sensors. Input is used. The 2D ASIC has four energy bins with a linear energy response over the dynamic range between 30 keV and 140 keV for clinical CT. A uniform energy resolution over the entire dynamic range indicates that charge collection from the direct conversion sensor is successful. Energy resolution is maintained at a high x-ray flux over varying long exposures to x-rays exhibiting polarization in a run unaffected by the sensor. The results demonstrate a fast output count rate from a clinical CT X-ray source with good energy resolution and low noise floor. Sensors and ASICs should be designed to accommodate existing clinical CT systems.

  Photon counting detectors using CdTe or CZT substrates are promising candidates for CT systems, but suffer from problems including charge sharing and pulse pile-up. Smaller micropixels, as well as minimizing increased pulse pileup, by using a hybrid pixel pattern design that incorporates increased micropixel size for the detector to improve charge sharing characteristics Thanks to the use of, there is no sacrifice of damage due to the increased pile-up. Hybrid pixel pattern design is an important design consideration in CdTe detectors. Using a hybrid pixel pattern design allows for better results in both high flux and low flux scanning environments without having to determine the optimal pixel size depending on the specific task and charge build time.

  In the hybrid pixel pattern design described in the present embodiment, when pile-up occurs in a large micropixel, the micropixel to be used is switched, such as using only a small micropixel without using a large micropixel. May be used to acquire the detection data.

  As a determination method for switching the micro pixel, for example, based on detection data at the time of capturing a positioning image (scano image), or to estimate a pile-up state in capturing of the next view, If the value of the detection data (such as the energy value) is equal to or greater than the threshold, the processing circuit 44 (eg, the system control function 441 or the pre-processing function 442) uses only small micropixels in capturing the next view. You may switch. Note that the size of the micropixel used in the shooting target view may be determined based on the value of the detection data of the two or more previous views, not limited to the immediately preceding view. The present invention is not limited thereto, and any general method for determining whether or not detector data piles up can be applied.

  The size of the micropixel to be used may be determined according to the position of the X-ray detector 12. For example, only small micropixels are used at the column and / or channel ends of the X-ray detector 12, which are expected to emit high flux X-rays, and low flux X-rays are emitted Both large and small micropixels may be used in the column and / or channel center of the X-ray detector 12, which is expected to be performed. The end portion and the center portion may be specified by a user, for example, or may be automatically determined by the processing circuit 44.

  Also, after the DAS 18 obtains the detection data from the PCD 12, the detection data obtained by the large micropixel is not used for the subsequent processing (such as generation of a spectral image), and only the detection data obtained by the small micropixel is used. You may. That is, the detection data obtained by the small micropixels may be left and the detection data obtained by the large micropixels may be switched and used.

  Whether the detection data of the large micropixel is discarded depends on whether the DAS 18 or the processing circuit 44 (for example, the system control function 441 or the preprocessing function 442) has piled up the detection energy detected by the large micropixel, or If the energy of the data is equal to or larger than the threshold, it may be determined not to use (discard) the detection data of the large micropixel. Alternatively, when the detection data is used, processing may be performed such as reducing the weight of the detection data of the large micropixel and increasing the weight of the detection data of the small micropixel.

  As described above, by using detection data of small micropixels that have not been piled up, reliability of the detection data can be improved and detection accuracy can be improved.

Conversely, in a low-flux scan environment, only the detection data of the large micropixels may be used without using the detection data of the small micropixels.
In this case, similarly to the switching of the PCD 12 described above, the micro pixel used in the PCD 12 may be switched, or after the detection data is acquired from the PCD 12, only the detection data of the large micro pixel is detected in the DAS 18 or the processing circuit 44. May be used. Alternatively, the DAS 18 or the processing circuit 44 may increase the weight of the detection data of the large micropixel and decrease the weight of the detection data of the small micropixel, and use the detection data in the subsequent processing.
Further, since the PCD 12 according to the present embodiment has two types of pixel patterns, the calibration of the PCD 12 becomes easier as compared with the case where three or more pixel pattern designs are used.

Here, the above-mentioned macro pixel and micro pixel are conceptual. A configuration example of the actually assumed X-ray detector 12 will be described with reference to FIG.
FIG. 6 is a diagram expressing the hybrid pixel pattern shown in FIG. 5C as an actual configuration example. Actually, an anode electrode 99 (also referred to as an anode pattern) made of a substance having a low electric resistance such as a metal is formed on the crystal 71 as a macro pixel 91 which is an example of the anode pixel 74. These are electrically insulated or arranged with discontinuous gaps 97.
With the configuration of the anode electrode 99 and the gap 97, a hybrid pixel pattern including the large micropixels 93 and the small micropixels 95 is formed.

Referring now to FIG. 7A, the various micropixels in the hybrid pixel pattern design include switches at the input stage 200 of the amplifier and counter. The channel providing the input to a given counter may change dynamically over the course of a CT scan or from scan to scan. This allows for configurable and dynamic management of the optimal and effective pixel size and the trade-off between pulse pile-up at the amplifier / counter stage and crosstalk at the detector pixels. Pixels P1, P2, and P3 may be of uniform size and shape, or may have a mixture of sizes and shapes instead. For example, at low count rates (less than 10 Mcps / mm ² ), pixel signals may be input together into the same amplifier / counter stage, thereby reducing the deleterious effects of crosstalk. Similarly, at high count rates of 100 Mcps / mm ^{2 and} higher, the pixel signal may be switched to a separate amplifier / counter stage to mitigate the deleterious effects of pulse pile-up. Box 202 represents an input channel from a PCD with a hybrid pixel pattern design. P1, P2, P3 in the example represent micropixels, which can include combinations of micropixels of different sizes. Each of the illustrated micropixels may be connected by a switch to a counter stage and amplifier indicated by box 206.

  In FIG. 7A, the micropixel P1 can be connected to the counter stage C1 and the amplifier via a switch S1a. Similarly, micropixel P2 can be connected to C2 via switch S2a, and micropixel P3 can be connected to C3 via switch S3. Alternatively, switch S1b can be connected to micropixels P1 and P2, and switch S2b can be connected to micropixels P2 and P3. In FIG. 7A, the three micropixels (P1, P2, P3) shown are connected together by switches S1b and S2b. In the illustrated example, the signals from pixels P1 and P2 are sent to counter C3. This is achieved by having switches S1a and S2a open while switches S1b and S2b are closed. If pixels P1, P2, P3 are physically adjacent, creating an effective pixel size of P1 + P2 + P3, which may result in an improvement in crosstalk performance, rather than separating pixels P1, P2, P3.

  FIG. 7B shows that the signal from pixel P1 is sent to counter C1 via switch S1a, the signal from pixel P2 is sent to counter C2 via S2a, and the signal from pixel P3 is sent to counter C3 via S3. It is another illustration example sent to. By keeping the pixels separate, pulse pile-up problems can be minimized and may be better in high flux scanning environments. In FIG. 7C, pixel P1 is sent to counter C1 via switch S1a, and pixels P2 and P3 are sent to C3 via switches S2b and S3. The channel providing the input to a particular counter may change dynamically to configure the optimal and effective pixel size for the hybrid pixel pattern associated with each macropixel of the photon counting detector. .

  Although several embodiments of the present invention have been described, these embodiments are provided by way of example and are not intended to limit the scope of the invention. These embodiments can be implemented in other various forms, and various omissions, replacements, and changes can be made without departing from the spirit of the invention. These embodiments and their modifications are included in the scope and gist of the invention, and are also included in the invention described in the claims and equivalents thereof.

DESCRIPTION OF SYMBOLS 1 X-ray CT apparatus 10 Mounting apparatus 11 X-ray tube 12 X-ray detector (PCD)
Reference Signs List 13 rotating frame 14 X-ray high voltage device 15 control device 16 wedge 17 collimator 18 data acquisition device (DAS)
19 Opening (bore)
Reference Signs List 30 bed apparatus 31 base 32 bed driving apparatus 33 top plate 34 support frame 40 console apparatus 41 memory 42 display 43 input interface 44 processing circuit 441 system control function 442 preprocessing function 443 reconstruction processing function 444 display control function

Claims (17)

A plurality of macropixels disposed on a semiconductor crystal having a first surface and a second surface parallel to the first surface, wherein each of the plurality of macropixels generates a reconstructed image; Collect projection data for
Each of the macro pixels is
Consisting of only two types: at least one large micro pixel arranged in each macro pixel and at least two small micro pixels arranged in each macro pixel;
The photon counting detector, wherein an area of each of the at least two small micropixels is smaller than an area of the at least one large micropixel.   In the macro pixel, half the length of one side of the macro pixel corresponds to at least one side of the large micro pixel, and half the length of one side of the large micro pixel is at least the length of the small micro pixel. The photon counting detector according to claim 1, which corresponds to a length of one side.   3. The photon counting detector of claim 1 or claim 2, wherein the at least one large micropixel has a pixel size that is an integer greater than one than the pixel size for the at least two small micropixels.   4. The photon counting detector according to claim 1, wherein the large micropixel and the small micropixel are similar. 5.   The photon counting detector according to claim 1, wherein the at least one large micropixel has a pixel size in a range from approximately 300 μm to approximately 600 μm.   The photon counting detector according to claim 1, wherein the at least two small micropixels have a pixel size in a range from approximately 150 μm to approximately 300 μm.   7. The photon counting detector according to any one of the preceding claims, wherein each macropixel comprises one large micropixel and 12 small micropixels.   8. The one large micro pixel of claim 7, wherein the one large micro pixel is disposed at a central portion of each macro pixel while an enclosure is formed around the one large micro pixel by the twelve small micro pixels. Photon counting detector.   7. The photon counting detector according to any one of the preceding claims, wherein each macropixel comprises two large micropixels and eight small micropixels.   7. A photon counting detector according to any one of the preceding claims, wherein each macropixel comprises three large micropixels and four small micropixels.   A photon counting detector according to any of the preceding claims, comprising a 16x16 array comprising 256 macropixels.   12. Photon counting detection according to any one of the preceding claims, wherein the pixel size of the at least one large micropixel is twice as large as the pixel size of one of the at least two small micropixels. vessel.   The photon counting detector of claim 1, comprising a first set of detector responses for the at least one large micropixel and a second set of detector responses for the at least two small micropixels.   14. The photon counting detector of claim 13, wherein information obtained from the first set of detector responses is used to calibrate the at least two small micropixels in a low flux scan environment.   15. A photon counting detector according to claim 13 or claim 14, wherein information obtained from the second set of detector responses is used to calibrate the at least one large micropixel in a high flux scanning environment. . A cathode electrode that covers the first surface;
The photon counting detector according to claim 1, wherein the plurality of macropixels cover the second surface. An X-ray tube for irradiating X-rays,
The photon counting detector according to any one of claims 1 to 15, wherein the photon counting detector detects X-rays emitted from the X-ray tube and transmitted through the subject.
An X-ray CT apparatus comprising: