Patent Application
20080165922
The present invention relates generally to diagnostic imaging and, more particularly, to a laminated CT detector collimator and methods of making same.
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 a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about a gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator having a plurality of collimator plates for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
An x-ray detector may, instead of using a scintillating device, include an energy discriminating detector having a direct conversion material capable of x-ray counting and capable of providing a measurement of the energy level of each x-ray detected. A laminated collimator as described herein is equally applicable to use with an energy discriminating device or other detector using pixellated elements.
Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Image quality can be directly associated with the degree of alignment between the components of the detector. “Cross-talk” between detector cells of a CT detector is common and to some degree is affected by the alignment, or lack thereof, of the detector components. In this regard, cross-talk is typically higher when the components of the CT detector are misaligned.
Cross-talk is generally defined as the communication of data between adjacent cells of a CT detector. Generally, cross-talk is sought to be reduced because cross-talk leads to artifact presence in the final reconstructed CT image and contributes to poor spatial resolution. Different types of cross-talk may result within a single CT detector. Cross-talk can occur as light from one cell is passed to another through a contiguous layer between the photodiode layer and the scintillator. Electrical cross-talk can occur from unwanted communication between photodiodes. Optical cross-talk may occur through the transmission of light through the reflectors that surround the scintillators. X-ray cross-talk may occur due to x-ray scattering between scintillator cells.
In an effort to reduce cross-talk, plates or layers of a collimator may be aligned with the cells of the scintillator arrays. The alignment of the cells of the scintillator arrays and the plates of the collimator can be a time consuming and labor intensive process. A collimator is typically fabricated using approximately 1000 collimating plates that are inserted between a set of rails. The rails typically have combs attached thereto, each comb having a plurality of teeth that are constructed to hold the collimating plates. Typically the rails are aligned to very exacting tolerances such that the teeth of the combs are positioned to receive the collimating plates, and, when inserted into the teeth, provide a collimating effect to the pixellating elements. Further, the physical placement or alignment of the collimator to the scintillator array is particularly susceptible to misalignment stack-up. That is, one of the scintillator-collimator assemblies, if unaligned, can detrimentally affect the alignment of adjacent assemblies. Simply, if one collimator-scintillator array combination is misaligned, all subsequently positioned collimator-scintillator array combinations will be misaligned absent implementation of corrective measures. Further, such assemblies require adjusting several detectors when only one of the detectors is misaligned. The overall process can be costly and time-consuming.
Mechanical deflection of the plates can occur due to G loading induced by rotation of the gantry at high gantry speeds. If a CT system is calibrated at, for instance, one speed and image data is acquired at, for instance, a second speed, then mechanical deflection of the collimator plates may be different for the two gantry speeds. Such mechanical deflection may induce image artifacts in resulting images. Additionally, Z-axis coverage of patients has increased in recent years and, and increased Z-axis coverage requires, likewise, proportionately longer collimator plates. Accordingly, the collimator plates are thus increasingly susceptible to mechanical deflection and the image quality problems resulting therefrom.
Therefore, it would be desirable to design a method and apparatus for the fabrication of a low-cost collimator and a scintillator module to thereby reduce cross-talk and increase mechanical stability thereof.
The present invention is directed to an apparatus that overcomes the aforementioned drawbacks. The CT detector includes a plurality of pixellated elements and a laminated collimator. Laminations within the collimator are separated by a spacer material and have apertures aligned between a respective pixellating element and an x-ray source.
According to one aspect of the present invention, a CT collimator includes a first radiation absorbent lamination having a plurality of apertures formed therethrough. Each aperture formed through the first radiation absorbent lamination is aligned with a respective axis formed between a corresponding pixellating element and an x-ray emission source. The collimator includes a second radiation absorbent lamination having a plurality of apertures formed therethrough, each aperture formed through the second radiation absorbent lamination aligned with the respective axis formed between a corresponding pixellating element and the x-ray emission source. A spacer is positioned between the first and second radiation absorbent laminations.
According to another aspect of the present invention, a method of fabricating a CT detector, includes providing a detector having a plurality of pixellated elements and coupling a multi-laminate collimator to the detector. The multi-laminate collimator includes at least two layers of material substantially impermeable to radiation. The method includes positioning an insert between the at least two layers, and aligning the collimator such that a plurality of x-ray passageways within the collimator are aligned between the plurality of pixellated elements and an x-ray emission source in a 1:1 correspondence.
In accordance with another aspect of the present invention, a CT system includes a rotatable gantry having an opening to receive an object to be scanned, a high frequency electromagnetic energy projection source configured to project a high frequency electromagnetic energy beam toward the object, and a detector array having a plurality of pixellated cells wherein each cell is configured to detect high frequency electromagnetic energy passing through the object. A radiation filter is configured to absorb high frequency electromagnetic energy directed toward a space between adjacent pixellated cells, wherein the radiation filter includes a pair of perforated screens separated at least by a spacer material. A photodiode array is optically coupled to the scintillator array and includes a plurality of photodiodes configured to detect light output from a corresponding scintillator cell. A data acquisition system (DAS) is connected to the photodiode array and configured to receive the photodiode outputs. An image reconstructor connected to the DAS and configured to reconstruct an image of the object from the photodiode outputs received by the DAS.
Various other features and advantages of the present invention will be made apparent from the following detailed description and the drawings.
The drawings illustrate one preferred embodiment presently contemplated for carrying out the invention.
In the drawings:
The operating environment of the present invention is described with respect to a sixty-four-slice computed tomography (CT) system. However, it will be appreciated by those skilled in the art that the present invention is equally applicable for use with other multi-slice configurations. Moreover, the present invention will be described with respect to the detection and conversion of x-rays. However, one skilled in the art will further appreciate that the present invention is equally applicable for the detection and conversion of other high frequency electromagnetic energy. The present invention will be described with respect to a “third generation” CT scanner, but is equally applicable with other CT systems.
Referring to
Rotation of gantry 12 and the operation of x-ray source 14 are governed by a control mechanism 26 of CT system 10. Control mechanism 26 includes an x-ray controller 28 that provides power and timing signals to an x-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of gantry 12. An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed reconstruction. The reconstructed image is applied as an input to a computer 36 which stores the image in a mass storage device 38.
Computer 36 also receives commands and scanning parameters from an operator via console 40 that has a keyboard. An associated cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from computer 36. The operator supplied commands and parameters are used by computer 36 to provide control signals and information to DAS 32, x-ray controller 28 and gantry motor controller 30. In addition, computer 36 operates a table motor controller 44 which controls a motorized table 46 to position patient 22 and gantry 12. Particularly, table 46 moves portions of patient 22 through a gantry opening 48.
As shown in
Detectors 20 include pins 52 positioned within pack 51 relative to detector elements 50. Pack 51 is positioned on diode array 53 having a plurality of diodes 59. Diode array 53 is in turn positioned on multi-layer substrate 54. Spacers 55 are positioned on multi-layer substrate 54. Detector elements 50 are optically coupled to diode array 53, and diode array 53 is in turn electrically coupled to multi-layer substrate 54. Flex circuits 56 are attached to face 57 of multi-layer substrate 54 and to DAS 32. Detectors 20 are positioned within detector assembly 18 by use of pins 52.
In operation, x-rays impinging within detector elements 50 generate photons which traverse pack 51, thereby generating an analog signal which is detected on a diode 58 within diode array 53. The analog signal generated is carried through multi-layer substrate 54, through one of flex circuits 56, to DAS 32 wherein the analog signal is converted to a digital signal.
A first laminate, or screen, 112 of collimator 110 is positioned proximate to scintillator pack 104. Laminate 112 is perforated such that each perforation, or aperture, 116 formed therein is sized and positioned to allow x-rays 16 to pass therethrough to impinge on an upper surface 120 of a corresponding pixel 106. In this manner, each perforation 116 is substantially aligned with an axis 137 formed between the corresponding pixel 106 and the focal spot 102. The perforations 116 of laminate 112 are further sized and positioned such that structural material 122 of laminate 112 is positioned to obstruct x-rays 16 that emit from focal spot 102 toward reflectors 108.
A second laminate, or screen, 114 of collimator 110 is positioned proximate to laminate 112 is perforated such that each perforation, or aperture 118 formed therein is sized and positioned to allow x-rays 16 to pass therethrough to impinge on an upper surface 120 of a corresponding pixel 106. In this manner, each perforation 118 is substantially aligned with an axis, one of which is illustrated as axis 137, formed between the corresponding pixel 106 and the focal spot 102. Accordingly, each pair of perforations 116 and 118 corresponding to each pixel 106 forms a hole, or opening, 129 through collimator 110 that is substantially aligned with a respective axis 137 formed between the corresponding pixel 106 and the focal spot 102. The perforations 118 of laminate 114 are further sized and positioned such that structural material 124 of laminate 114 is positioned to obstruct x-rays 16 that emit from focal spot 102 toward reflectors 108.
A fanout angle 128 is formed between one pixel 106 and focal spot 102 of one axis 139 and another fanout angle 130 is formed between another pixel 106 and focal spot 102 of another axis 137. A pattern of the perforations 116 of the first laminate 112 may be distinct from a pattern of the perforations 118 of the second laminate 114. Accordingly, in one embodiment, perforations 116 have a larger opening and are positioned closer together than respective perforations 118. In another embodiment, perforations 116 are positioned further apart than respective perforations 118 but have an opening substantially similar to an opening of respective perforations 118. It is contemplated, however, that the patterns of respective perforations 116, 118 are substantially similar and are sized and positioned according to the fanout as defined by, for instance, fanout angles 128 and 130. Additionally, according to the fanout angle 128, perforations 116 in laminate 112 may have different sizes and spacings with respect to each other. Similarly, according to the fanout angle 130, perforations 118 in laminate 118 may have different sizes and spacings with respect to each other.
Laminae 112, 114 comprise a high density material, such as tungsten or the like. Accordingly, laminae 112, 114 are substantially impermeable to and substantially attenuate x-rays 16 that would otherwise impinge on the region of reflectors 108 in scintillator pack 104. It is contemplated that the perforations 116, 118 therein are fabricated by etching, drilling, molding, or the like.
Still referring to
In operation, collimator 110 substantially attenuates x-rays 16 that emit from focal spot 102 from impinging on reflectors 108. Collimator 110 also collimates x-rays that emit from a secondary emission point 133 within, for instance, patient 22 of
Referring still to
Furthermore, while CT detector 100 is illustrated in
Referring again to
Referring to
Referring now to
Therefore, according to one embodiment of the present invention, a CT collimator includes a first radiation absorbent lamination having a plurality of apertures formed therethrough. Each aperture formed through the first radiation absorbent lamination is aligned with a respective axis formed between a corresponding pixellating element and an x-ray emission source. The collimator includes a second radiation absorbent lamination having a plurality of apertures formed therethrough, each aperture formed through the second radiation absorbent lamination aligned with the respective axis formed between a corresponding pixellating element and the x-ray emission source. A spacer is positioned between the first and second radiation absorbent laminations.
According to another embodiment of the present invention, a method of fabricating a CT detector, includes providing a detector having a plurality of pixellated elements and coupling a multi-laminate collimator to the detector. The multi-laminate collimator includes at least two layers of material substantially impermeable to radiation. The method includes positioning an insert between the at least two layers, and aligning the collimator such that a plurality of x-ray passageways within the collimator are aligned between the plurality of pixellated elements and an x-ray emission source in a 1:1 correspondence.
In accordance with another embodiment of the present invention, a CT system includes a rotatable gantry having an opening to receive an object to be scanned, a high frequency electromagnetic energy projection source configured to project a high frequency electromagnetic energy beam toward the object, and a detector array having a plurality of pixellated cells wherein each cell is configured to detect high frequency electromagnetic energy passing through the object. A radiation filter is configured to absorb high frequency electromagnetic energy directed toward a space between adjacent pixellated cells, wherein the radiation filter includes a pair of perforated screens separated at least by a spacer material. A photodiode array is optically coupled to the scintillator array and includes a plurality of photodiodes configured to detect light output from a corresponding scintillator cell. A data acquisition system (DAS) is connected to the photodiode array and configured to receive the photodiode outputs. An image reconstructor connected to the DAS and configured to reconstruct an image of the object from the photodiode outputs received by the DAS.
The present invention has been described in terms of the preferred embodiment, and it is recognized that equivalents, alternatives, and modifications, aside from those expressly stated, are possible and within the scope of the appending claims.
