JP2008224302A - Focally-aligned ct detector - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To design a CT detector in which artifacts induced by abnormal alignment is less likely to occur and provide a CT detector having detector cells consistently aligned with an X-ray source of a radiation imaging system. <P>SOLUTION: A focally aligned scintillator 57 is constructed, such that its scintillator walls 90 are sloped so as to be aligned in angular manner with the X-ray source 14. The scintillator 57 has a planar X-ray receiving surface 86 and a planar light emitting surface 88, and a plurality of sidewalls 90 connecting a planar X-ray receiving surface 86 and the planar light emitting surface 88. The sidewalls 90 extend non-perpendicularly between the planar X-ray reception surface 86 and the planar light emitting surface 88. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention relates generally to diagnostic imaging, and more specifically to a radiation imaging detector having a focus aligned cell.

  Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or baggage. In the following, the terms “subject” and “subject” shall include any object that can be imaged. The beam enters the array of radiation detectors after being attenuated by the subject. The intensity of the attenuated beam radiation received by the detector array typically depends on the attenuation of the x-ray beam by the subject. Each detector element of the detector array generates a separate electrical signal that indicates the attenuated beam received by each detector element. The electrical signal is transmitted to the data processing system and analyzed, from which it ultimately forms an image.

  In general, the x-ray source and detector array rotate around the gantry around the subject in the imaging plane. The x-ray source typically includes an x-ray tube that emits an x-ray beam at the focal point. An X-ray detector typically includes a collimator that collimates an X-ray beam received by the detector, a scintillator that is provided adjacent to the collimator and converts X-rays into optical energy, and an adjacent scintillator. And a photodiode that receives light energy and generates an electrical signal therefrom.

  Typically, each scintillator in the scintillator array converts x-rays into light energy. Each scintillator emits light energy toward a photodiode adjacent to the scintillator. Each photodiode detects light energy and generates a corresponding electrical signal. The output of the photodiode is then transmitted to the data processing system for image reconstruction.

  Despite many advances being made for known CT detectors, image quality is still an important area and is an area that needs improvement. Specifically, there is still a need to improve image quality by reducing image artifacts. Although image artifacts can be attributed to a number of factors, one problem faced by conventional CT detectors is misalignment of the scintillator with respect to the x-ray source or post-patient collimator. The adverse effects of misaligned scintillators are shown in FIG.

  FIG. 8 is a cross-sectional view of a conventional CT detector 2. The detector includes a scintillator array 4 of scintillators 6. The scintillator array is positioned above the photodiode array (not shown), and the light emitted by the scintillator array in response to receiving X-rays 7 is detected and processed by the photodiode array. It is like that. For illustrative purposes, the scintillator array also includes a single misaligned scintillator 6 (a). Conventional detector designs also include an x-ray shielding element 8. These shielding elements are designed to block X-rays and, as a result, typically block some X-rays 7 (a) from passing through the inter-cell gap 9, The X-ray 7 (b) cannot be shut off and is allowed to pass through the inter-cell gap 9.

  Due to misalignment of the scintillator with respect to the X-ray source, the X-rays pass through various thicknesses of scintillator material, resulting in spectral gain inhomogeneities in the scintillator, such as bone-induced spectral artifacts. That is, in a scintillator that is not aligned with other scintillators, the X-ray path length is different from other scintillators. Thus, these misaligned scintillators have a different response than a correctly aligned scintillator with respect to the spectrum.

  In other words, the path length of X-rays passing through the scintillator gap is different from that of X-rays passing only through the scintillator. Due to this path length difference, misaligned scintillators have different responses to adjacent correctly aligned scintillators. In addition, misaligned shielding elements can also contribute to spectral non-linearities that occur when the scintillator is misaligned. As a result, conventional CT detectors are susceptible to artifacts induced by misalignment of detector cells, such as ring, band and center artifacts.

  Therefore, it is desirable to design a CT detector that is less prone to artifacts induced by misalignment. It would further be desirable to provide a CT detector having a detector cell that is consistently aligned with the x-ray source of the radiation imaging system.

  The present invention relates to a focus-aligned CT detector that overcomes the above-mentioned drawbacks. The CT detector is constructed by tilting the scintillator wall so that it is aligned with respect to the X-ray source with respect to angle. In this regard, the CT detector is less prone to spectral artifacts associated with detector cell misalignment.

  Thus, according to one aspect, the present invention includes a scintillator having a planar X-ray receiving surface and a planar light emitting surface. The scintillator also has a plurality of sidewalls connecting the planar X-ray receiving surface and the planar light emitting surface. The sidewall extends non-vertically between the planar X-ray receiving surface and the planar light emitting surface.

  According to another aspect of the invention, a radiation detector includes a photodiode array including a plurality of photodiodes configured to output an electrical signal in response to sensed light. Each photodiode has a planar light detection surface. The detector further includes a scintillator array including a plurality of scintillators configured to emit light in response to receiving X-rays. Each scintillator has sidewalls that are inclined with respect to the planar photodetection surface of the respective photodiode.

  According to another aspect, the present invention includes a CT system having a rotating gantry. The gantry has an opening that accommodates the object to be scanned. The system also includes an x-ray source configured to project an x-ray fan beam at a given projection angle toward the subject and a plurality of scintillators configured to convert x-ray energy into light. A scintillator array having cells; Each scintillator cell is defined by an off-centered sidewall that extends along an angle parallel to a given projection angle. A photodiode array is optically coupled to the scintillator array and includes a plurality of photodiodes configured to detect light emitted from the scintillator array and provide an electrical signal output. . The system further includes a data acquisition system (DAS) connected to the photodiode array and configured to receive an electrical signal output of the photodiode array, and a photodiode connected to the DAS and received by the DAS. An image reconstructor configured to reconstruct an image of interest from the array electrical signal output;

  Various other features and advantages of the present invention will be made apparent from the following detailed description and the drawings.

  The drawings illustrate a preferred embodiment presently contemplated for carrying out the invention.

  The operating environment of the present invention will be described with respect to a 4-slice computed tomography (CT) system. However, those skilled in the art will recognize that the present invention is equally applicable for use in single-slice configurations or other multi-slice configurations. Further, the present invention will be described with respect to X-ray detection and conversion. However, those skilled in the art will further appreciate that the present invention is equally applicable to the detection and conversion of other high frequency electromagnetic energy. Also, although the present invention will be described with respect to a “third generation” CT scanner, the present invention is equally applicable to other CT systems. It is also contemplated that the present invention is applicable to detectors in other radiation imaging systems such as X-ray scanners.

  1 and 2, a computed tomography (CT) imaging system 10 is shown as including a gantry 12 typical of a “third generation” CT scanner. The gantry 12 includes an X-ray source 14 that projects an X-ray beam 16 toward a detector array 18 provided on the opposite side of the gantry 12. The detector array 18 is formed by a plurality of detectors 20, and the detectors 20 collectively detect the projected X-rays transmitted through the patient 22. Each detector 20 generates an electrical signal that represents the intensity of the incident x-ray beam and thus represents the attenuated beam as it passes through the patient 22. The gantry 12 and components mounted on the gantry 12 rotate around the rotation center 24 during one scan for acquiring X-ray projection data.

  The rotation of the gantry 12 and the operation of the X-ray source 14 are controlled by the control mechanism 26 of the CT system 10. The control mechanism 26 includes an X-ray controller 28 and a gantry motor controller 30. The X-ray controller 28 supplies power signals and timing signals to the X-ray source 14, and the gantry motor controller 30 The rotational speed and position of the gantry 12 are controlled. A data acquisition system (DAS) 32 provided in the control mechanism 26 samples the analog data from the detector 20 and converts this data into a digital signal for subsequent processing. An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed image reconstruction. The reconstructed image is applied as an input to the computer 36, which causes the mass storage device 38 to store the image.

  The computer 36 also receives commands and scanning parameters from an operator via a console 40 having a keyboard. The attached cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from the computer 36. The commands and parameters supplied by the operator are used by the computer 36 to supply control signals and information to the DAS 32, X-ray controller 28 and gantry motor controller 30. In addition, the computer 36 operates the table motor controller 44 that controls the motorized table 46 to place the patient 22 and the gantry 12. Specifically, the table 46 moves each part of the patient 22 through the gantry opening 48.

  As shown in FIGS. 3 and 4, the detector array 18 includes a plurality of scintillators 57 that form a scintillator array 56. A post patient collimator (not shown) is placed above the scintillator array 56 to collimate the X-ray beam 16 before it enters the scintillator array 56.

  In one embodiment shown in FIG. 3, the detector array 18 includes 57 detectors 20, each detector 20 having a 16 × 16 array size. As a result, the array 18 has 16 rows and 912 columns (16 × 57 detectors) and can collect 16 slices of data simultaneously with each rotation of the gantry 12.

  Switch arrays 80 and 82 in FIG. 4 are multidimensional semiconductor arrays coupled between scintillator array 56 and DAS 32. Switch arrays 80 and 82 include a plurality of field effect transistors (FETs, not shown) arranged as a multidimensional array. The FET array includes a fixed number of electrical leads connected to each of the respective photodiodes 60 and a fixed number of output leads electrically connected to DAS 32 via a flexible electrical interface 84. It is out. Specifically, about half of the photodiode output is electrically connected to the switch 80, and the other half of the photodiode output is electrically connected to the switch 82. In addition, a reflector layer (not shown) may be interposed between each scintillator 57 to reduce light scattering from adjacent scintillators. Each detector 20 is fixed to the detector frame 77 of FIG.

  Switch arrays 80 and 82 further include a decoder (not shown) that enables, disables and combines the photodiode outputs according to the desired number of slices and the desired slice resolution for each slice. The decoder is in one embodiment a decoder chip or FET controller known in the art. The decoder includes switch arrays 80 and 82 and a plurality of output and control lines coupled to DAS 32. In one embodiment defined as 16 slice mode, the decoder enables switch arrays 80 and 82 to activate all rows of photodiode array 52, resulting in 16 slices being processed by DAS 32. Minute data is generated at the same time. Needless to say, many other slice combinations are possible. For example, the decoder may select from other slice modes including 1 slice mode, 2 slice mode, and 4 slice mode.

  As shown in FIG. 5, the switch arrays 80 and 82 are configured in a 4-slice mode by sending the appropriate decoder instructions so that the data consists of one or more rows of the photodiode array 52. It can be collected from a single slice. Depending on the specific configuration of the switch arrays 80 and 82, various combinations of photodiodes 60 are enabled so that the slice thickness can consist of one, two, three or four rows of scintillator array elements 57. Can be disabled or combined. Other examples include single slice mode with one slice from 1.25 mm to 20 mm thick slice, and two slice mode with two slices from 1.25 to 10 mm thick slices. There is. Other modes are also conceivable.

  FIG. 6 shows a cross-sectional view of a CT detector 20 according to the present invention. For illustration purposes, only five scintillators and photodiodes are shown, but those skilled in the art will appreciate that a CT detector can include more such scintillators and photodiodes. Also, as is well known, scintillator arrays and photodiode arrays are two-dimensional (2D) arrays. As illustrated and described above, the detector 20 includes a scintillator array 56 comprised of a plurality of scintillators 57 that emit light upon receipt of X-ray energy. This emission is detected by the photodiode 60 of the photodiode array 52. In this regard, each scintillator 57 has a planar X-ray receiving surface 86 and a planar light emitting surface 88. The surfaces 86, 88 are connected to each other by scintillator partitions or sidewalls 90. As shown, the sidewall 90 is angled with respect to the x-ray receiving surface and the light emitting surface. As a result of the inclination angle of the sidewall, the X-ray receiving surface 86 of the scintillator 57 is displaced from the light emitting surface 88 of the scintillator 57.

  The inclined sidewall 90 is angled so that the scintillator is focused on an x-ray source (not shown). In this respect, the side wall is inclined parallel to the X-ray path 16. By this inclination, the side walls are oriented non-perpendicular to the X-ray receiving surface and the light emitting surface. Further, the partition wall is angled with respect to the surface 91 of the photodiode 52. As a result, the x-ray path is relatively uniform and constant between the scintillators of the scintillator array. This is particularly advantageous for misaligned scintillators such as those shown by scintillator 57 (a). In other words, the spectral response is less sensitive to scintillator misalignment because there is less path length variation for misaligned scintillators.

  With continued reference to FIG. 6, the CT detector 20 preferably includes a collimator 92 formed collectively by an array 94 of collimator elements or collimator plates. Preferably, each collimator plate is configured as an extension of the respective scintillator sidewall. Thus, like the scintillator array, the collimator grating is aligned with the x-ray source. In addition, it is contemplated that the detector 20 may be constructed with a shielding element (not shown) that provides additional x-ray collimation and x-ray separation. Furthermore, it is preferable to fill the scintillator gap 90 with a light reflective epoxy or other material to reduce optical crosstalk between the scintillators. The collimator plate 94 collectively forms a one-dimensional (1D) collimator 92.

  It is contemplated that the construction of the scintillator described above can be accomplished by one of a number of manufacturing techniques or a combination thereof. In this regard, the scintillator can be formed by casting a scintillator material. Alternatively, conventional molding techniques may be used. In addition, it is contemplated that mechanical or chemical cutting techniques may be used. Further, it is contemplated that electromagnetic ablation such as with a laser may be used. Regardless of the manufacturing technique, the scintillator wall is constructed to tilt towards the x-ray source of the radiation imaging system when placed in the detector assembly. As an advantage, this prevents the sidewall itself from being exposed to primary radiation during data acquisition.

  In FIG. 7, a parcel / parcel / baggage inspection system 100 includes a rotating gantry 102 having an opening 104 therein through which a parcel, parcel or baggage can pass. The rotating gantry 102 houses a high frequency electromagnetic energy source 106 and a detector assembly 108 having a scintillator array comprised of scintillator cells similar to those described above. A conveyor system 110 is also provided, which is supported by the structure 114 and automatically and continuously passes a parcel or baggage 116 through the opening 104 for scanning. 112 is included. The object 116 is fed into the opening 104 by the conveyor belt 112, and then imaging data is acquired and the parcel 116 is removed from the opening 104 by the conveyor belt 112 in a controlled, continuous manner. As a result, postal inspectors, baggage workers and other security personnel can non-invasively inspect the contents of parcels 116 for explosives, blades, guns, smuggled goods, and the like.

  Thus, according to one aspect, the present invention includes a scintillator having a planar X-ray receiving surface and a planar light emitting surface. The scintillator also has a plurality of sidewalls connecting the planar X-ray receiving surface and the planar light emitting surface. The sidewall extends non-vertically between the planar x-ray receiving surface and the planar light emitting surface.

  According to another aspect of the invention, a radiation detector includes a photodiode array including a plurality of photodiodes configured to output an electrical signal in response to sensed light. Each photodiode has a planar light detection surface. The detector further includes a scintillator array including a plurality of scintillators configured to emit light in response to receiving X-rays. Each scintillator has a sidewall that is inclined with respect to the planar photodetection surface of the respective photodiode.

  According to another aspect, the present invention includes a CT system having a rotating gantry. The gantry has an opening that accommodates the object to be scanned. The system also includes an x-ray source configured to project an x-ray fan beam at a given projection angle toward the subject and a plurality of scintillators configured to convert x-ray energy into light. A scintillator array having cells; Each scintillator cell is defined by an eccentric side wall that extends along an angle parallel to a given projection angle. A photodiode array is optically coupled to the scintillator array and includes a plurality of photodiodes configured to detect light emitted from the scintillator array and provide an electrical signal output. . The system further includes a data acquisition system (DAS) connected to the photodiode array and configured to receive an electrical signal output of the photodiode array, and a photodiode connected to the DAS and received by the DAS. An image reconstructor configured to reconstruct an image of interest from the array electrical signal output;

  While the invention has been described in terms of preferred embodiments, it will be appreciated that equivalent constructions, alternative constructions and modifications other than those explicitly described are possible and fall within the scope of the appended claims.

1 is a sketch of a CT imaging system. It is a block schematic diagram of the system shown in FIG. FIG. 6 is a perspective view of one embodiment of a CT system detector array. FIG. 6 is a perspective view of one embodiment of a detector. FIG. 5 is a diagram of various configurations of the detector of FIG. 4 in 4-slice mode. 2 is a partial cross-sectional view of a CT detector according to the present invention. FIG. 1 is a sketch of a CT system used with a non-invasive parcel inspection system. It is sectional drawing of the conventional CT detector.

Explanation of symbols

2 conventional CT detector 4 scintillator array 6 scintillator 7 X-ray 8 X-ray shielding element 9 inter-element gap 10 computed tomography (CT) imaging system 12 gantry 14 X-ray source 16 X-ray beam 18 detector Array 20 Multiple detectors 22 Patient 24 Center of rotation 26 Control mechanism 28 X-ray controller 30 Gantry motor controller 32 Data acquisition system (DAS)
34 Image Reconstructor 36 Computer 38 Mass Storage 40 Operator via Console 42 Cathode Ray Tube Display 44 Table Motor Controller 46 Motorized Table 48 Gantry Opening 52 Photodiode Array 56 Scintillator Array 57 Multiple scintillators 60 Photodiodes 77 Detector frame 79 Mounting brackets 80 and 82 Switch array 84 Flexible electrical interface 86 Planar X-ray receiving surface 88 Planar light emitting surface 90 Side wall 91 Surface 92 Collimator 94 Array of collimator elements or collimator plates 100 Parcel / parcel / baggage inspection system 102 Rotating gantry 104 Opening 106 High frequency electromagnetic energy source 108 Detector assembly 110 Conveyor system 11 Conveyor belt 114 structure 116 packages or pieces of baggage

Claims (9)

A planar X-ray receiving surface (86) and a planar light emitting surface (88);
The planar X-ray receiving surface (86) and the planar light emitting surface (88) are connected, and the planar X-ray receiving surface (86) and the planar light emitting surface (88) are connected. A plurality of side walls (90) extending non-vertically between
A scintillator (57) comprising:   The side wall (90) includes the planar x-ray receiving surface (86) and the planar light emitting surface (88) such that the side wall (90) is aligned with the x-ray source (14) during radiation imaging. The scintillator (57) according to claim 1, wherein the scintillator (57) is disposed at an angle to each other.   The scintillator (57) of claim 2, wherein the planar x-ray receiving surface (86) is linearly offset relative to the planar light emitting surface (88).   The scintillator (57) of claim 1, formed by casting a scintillator material.   The scintillator (57) of claim 1, formed by molding a scintillator material.   The scintillator (57) of claim 1, formed by cutting a scintillator bulk.   The scintillator (57) of claim 1, formed by electromagnetic ablation using a laser.   The scintillator (57) of claim 1, incorporated in a detector assembly (20) of a computed tomography (CT) scanner (10).   The scintillator (57) of claim 1, wherein the x-ray receiving surface (86) has a surface area equal to the light emitting surface (88).

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