Abstract:
One embodiment of the present invention is a method for measuring a radiation dose of a patient applied by a computerized tomography (CT) imaging system having a radiation source emitting a radiation beam and a radiation detector. The method includes steps of: scanning a patient with the radiation beam; imaging the patient utilizing radiation detected by the radiation detector; and estimating a patient radiation dose from said imaging utilizing a measurement of radiation delivered to only a portion of the detector during said imaging. 
     The present invention permits the advantages of a post-patient collimated detector array to be realized while providing an accurate estimate of patient radiation dosage.

Description:
BACKGROUND OF THE INVENTION 
     This invention relates generally to computed tomography (CT) imaging systems, and more particularly to methods and apparatus for measuring radiation dosages in post-patient collimated CT imaging systems. 
     In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile. 
     In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a “view”. A “scan” of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called “CT numbers” or “Hounsfield units”, which are used to control the brightness of a corresponding pixel on a cathode ray tube display. 
     At least one known CT imaging system employs post-patient collimation, one purpose of which is to reduce slice broadening. For example, a mechanical collimator is interposed between the patient and a radiation detector matrix. The mechanical collimator is comprised of a strip of material essentially opaque to x-rays, such as lead. The strip covers a detector assembly housing a detector array, and has a slot aligned along the detector elements to collimate x-rays impinging on the detector elements. 
     In imaging systems without post-patient collimation, an x-ray dose delivered to a patient can be directly measured from a measurement of the x-rays reaching the detector. However, in systems employing post-patient collimation, not all of the x-ray beam actually reaches the detector during dose testing, because some of it is blocked by the post-patient collimator. As a result, patient dosage cannot be accurately determined, because an unknown percentage of the radiation dose received by a patient never reaches the detection matrix. 
     It would therefore be desirable if methods and apparatus were available to provide accurate dose measurement even when post-patient collimation is employed. 
     BRIEF SUMMARY OF THE INVENTION 
     There is therefore provided, in one embodiment of the present invention, a method for measuring a radiation dose of a patient applied by a computerized tomography (CT) imaging system having a radiation source emitting a radiation beam and a radiation detector. The method includes steps of: scanning a patient with the radiation beam; imaging the patient utilizing radiation detected by the radiation detector; and estimating a patient radiation dose from said imaging utilizing a measurement of radiation delivered to only a portion of the detector during said imaging. 
     Utilizing methods and apparatus of the present invention, the advantages of a post-patient collimated detector array can be realized. At the same time, an accurate estimate of patient radiation dosage can be determined. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a pictorial view of a CT imaging system. 
     FIG. 2 is a block schematic diagram of the system illustrated in FIG.  1 . 
     FIG. 3 is a perspective view of a CT system detector array. 
     FIG. 4 is a perspective view of a detector module shown in FIG.  3 . 
     FIG. 5 is a perspective view of a portion of a post-patient collimator used in conjunction with the CT system detector array of FIG.  3 . 
     FIG. 6 is a simplified cross-sectional view of an x-ray beam directed at the detector array of FIG. 3 as used in the CT imaging system of FIG.  1 . 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring to FIGS. 1 and 2, a computed tomograph (CT) imaging system  10  is shown as including a gantry  12  representative of a “third generation” CT scanner. Gantry  12  has an x-ray source  14  that projects a beam of x-rays  16  toward a detector array  18  on the opposite side of gantry  12 . Detector array  18  is formed, in part, by detector elements  20  which together sense the projected x-rays that pass through an object  22 , for example a medical patient. Detector array  18  may be fabricated in a single slice or multi-slice configuration. Each detector element  20  produces an electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuation of the beam as it passes through patient  22 . During a scan to acquire x-ray projection data, gantry  12  and the components mounted thereon rotate about a center of rotation  24 . The rotation of gantry  12  also defines a z-axis perpendicular to a plane of gantry  12 . Linear dimensions parallel to the z-axis are denoted herein as “thicknesses.” 
     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 x-ray source  14  and a gantry motor controller  30  that controls the rotational speed and position of gantry  12 . A data acquisition system (DAS)  32  in control mechanism  26  samples analog data from detector elements  20  and converts the data to digital signals 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 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  in gantry  12 . Particularly, table  46  moves portions of patient  22  through gantry opening  48 . 
     As shown in FIGS. 3 and 4, detector array  18  includes a plurality of detector or imaging modules  20  that together form a imaging portion or region  50  of detector array  18 . In one embodiment, each detector module  20  includes a high-density semiconductor array  52  and a multidimensional scintillator array  54  positioned above and adjacent to semiconductor array  52 . Detector array  18  also has at least one additional measuring module  56  positioned adjacent at least one end  58  of detector array  18 . As used herein, “adjacent a longitudinal end” means touching an end  58 , or slightly separated therefrom but sufficiently close thereto so as to intercept a portion of x-ray beam  16 . For example, detector array  18  comprises measuring modules  56  touching both opposite longitudinal ends  58  of detector array  18 . Measuring modules  56  form measuring portions or regions  60  of detector array  18 . Measuring regions  60  are somewhat thicker in z-axis extent than imaging portion  50 . In one embodiment, this difference in thickness is a produced by using different size detector modules  20  for measuring region  60  and imaging portion  50  of detector array  18 . A post-patient collimator (not shown in FIGS. 3 or  4 ) is used to provide beam collimation of x-ray beam  16  to a thickness less than or equal to that of imaging portion  50 , while allowing a full measuring region  60  thickness of x-ray beam  16  to strike measuring region  60 . 
     In one embodiment, imaging portion  50  and measuring region  60  are the same thickness, and a post-patient collimator  62  (a portion of which is shown in FIG. 5) providing both scatter collimation and beam thickness collimation is positioned above and adjacent imaging portion  50  of scintillator array  54 . One portion  84  of post-patient collimator  62  provides both types of collimation for imaging portion  50  of scintillator array  54 . Collimator  62  provides scatter collimation with slots  72 , and beam collimation with rails  86 , which are thicker over imaging portion  50  of scintillator array  54  than over measuring portions  60 . Due to the opaqueness of the rails  86  blocking out a portion of x-ray beam  16  over detector array  18 , post-patient collimator  62  provides beam collimation of x-ray beams  16  before such beams impinge upon scintillator array  54 . A portion  88  of post-patient collimator  62  has thinner rails  86  and therefore does not provide beam collimation. 
     In one embodiment, x-ray beam  16  is subjected to post-patient beam collimation so that less than a thickness of beam  16  contributing to a radiation dose to patient  22  impinges imaging portion  50  of scintillator array  54 . As a result, a measurement of radiation dose cannot be made solely from radiation striking imaging portion  50 . However, the post-patient beam collimation is not applied in a manner that affects a peripheral portion of x-ray beam  16 . Therefore, a full thickness of x-ray  10  beam  16  impinges measuring portions  60  of scintillator array  54  and this full thickness is used to estimate the radiation dose of patient  22  during imaging. 
     “A full thickness of x-ray beam  16 ,” as used herein refers to a thickness encompassing essentially all of the radiation energy of x-ray beam  16  in a direction parallel to the z-axis. Because post-patient collimation of x-ray beam  16  blocks a portion of the energy of x-ray beam  16 , a measurement of a full thickness of x-ray beam  16  is more representative of a total patient  22  than would be a measurement of the post-patient beam collimated portion. In one embodiment, more accurate estimations of patient  22  radiation dosage are provided by configuring imaging system  10  being configured so that non-post patient collimated portions of x-ray beam  16  are neither attenuated by patient  22  nor significantly attenuated by portions of imaging system  10  before striking measuring portions  60 . 
     Semiconductor array  52  includes a plurality of photodiodes  64 , a switch apparatus  66 , and a decoder  68 . Photodiodes  64  are, for example, individual photodiodes. A multidimensional diode array construction is also acceptable. Photodiodes  64  are deposited or formed on a substrate (not shown). Scintillator array  54 , as known in the art, is positioned over and adjacent photodiodes  64 . Photodiodes  64  are optically coupled to scintillator array  54  and have electrical output lines  70  for transmitting signals representative of the light output by scintillator array  54 . Each photodiode  64  produces a separate low level analog output signal that is a measurement of beam attenuation for a specific scintillator of scintillator array  54 . Photodiode output lines  70  are, for example, physically located on one side of module  20  or on a plurality of sides of module  20 . In the embodiment illustrated in FIG. 4, photodiode outputs  70  are located at opposing sides of the photodiode array. 
     In one embodiment and as shown in FIG. 3, detector array  18  includes fifty-seven detector or imaging modules  20 . Each detector module  20  includes a semiconductor array  52  and scintillator array  54 , each having an array size of 16×16. As a result, array  18  comprises sufficient detector modules  20  for 16 rows and 912 columns of imaging (16×57 modules), thus allowing up to N=16 simultaneous slices of data to be collected along a z-axis with each rotation of gantry  12 . In another embodiment, detector modules  20  are single-slice detector modules having an array size of 1 row by 16 columns, so that array  18  has sufficient detector modules  20  for 1 row and 912 columns. However, the invention is not limited to a specific number of rows or slice thickness, or to a specific array size or number of columns. Slots  72  of post-patient collimator  62  are aligned over each column of detector modules  20  to collimate x-rays of x-ray beam  16  that impinge thereon. 
     Measuring module  56  is sufficiently thick to detect an entire thickness of x-ray beam  16  as measured in a direction along the z-axis. In other words, an effective thickness of measuring module  56 , in one embodiment, is as thick or thicker than a thickness of all slices measurable by measuring portion  50  and as thick as that portion of x-ray beam  16  having sufficient intensity to contribute to a patient radiation dose. In one embodiment, detection elements of measuring module  56  are scintillators and photodiodes similar to those of imaging modules  20 . However, measuring module  56  need not be configured for multiple slice operation nor have the same size or number of detection elements as imaging module  58 . In one embodiment, measuring module  56  is segmented similarly to imaging module  20 , allowing all or portions of its measuring region to be used for dose measurement when impinged by a partially obstructed x-ray beam. 
     In one multi-slice detector array embodiment, switch apparatus  66  is a multidimensional semiconductor switch array of similar size as that portion of semiconductor array  52  that includes detector or imaging modules  20 . Switch apparatus  66  is coupled between semiconductor array  52  and DAS  32 . Semiconductor device  66 , in one embodiment, includes two semiconductor switches  74  and  76 . Switches  74  and  76  each include a plurality of field effect transistors (FETs) (not shown) arranged as a multidimensional array. Each FET includes an input line electrically connected to one of the respective photodiode output lines  70 , an output line, and a control line (not shown). FET output and control lines are electrically connected to DAS  32  via a flexible electrical cable  78 . Particularly, about one-half of photodiode output lines  70  are electrically connected to each FET input line of switch  74  with the other one-half of photodiode output lines  70  electrically connected to FET input lines of switch  76 . Decoder  68  controls the operation of switch apparatus  66  to enable, disable, or combine photodiode outputs in accordance with a desired number of slices and slice resolutions for each slice. 
     Decoder  68  is a decoder chip or an FET controller as known in the art. Decoder  68  includes a plurality of output and control lines coupled to switch apparatus  66  and DAS  32 . Particularly, the decoder outputs are electrically coupled to the switch apparatus control lines to enable switch apparatus  66  to transmit the proper data from the switch apparatus inputs to the switch apparatus outputs. Utilizing decoder  68 , specific FETs within switch apparatus  66  are enabled, disabled, or combined so that specific photodiode outputs are electrically connected to CT system DAS  32  to provide selected slice thicknesses and/or a selected number of image slices. 
     Outputs of measuring modules  56  are also electrically connected to CT system DAS  32 , but data obtained from these outputs are not processed as image data. Instead, in one embodiment, DAS  32  and image reconstructor  34  recognize these inputs as providing x-ray dose data and pass this data on to computer  36 , which calculates an x-ray dosage based upon data received from measuring modules  56 . 
     In one embodiment, imaging system  10  scans patient  22  with radiation beam  16  and images patient  22  utilizing radiation detected by detector array  18 . 
     Patient radiation dosage is estimated from a measurement of radiation delivered to only a portion  60  of detector array  18  during imaging. 
     In one embodiment and referring to FIG. 6, x-ray source  14  is configured to project an x-ray beam  16  towards detector array  18 , which includes a beam collimated, imaging portion  50  and measuring portions  60  that are not beam collimated. X-ray source  14  and detector array  18  are mounted on a rotating gantry  12 , which is not shown in the simplified view of FIG.  6 . Table or patient support  46  supports patient  22  between x-ray source  14  and detector array  18  so that at least a portion  80  of x-ray beam  16  passes through patient  22  and is thereby attenuated. The attenuated radiation beam  80  is beam collimated to reduce its thickness and impinges on imaging portion  50 . Data representative of images of patient  22  are collected therefrom. 
     Another portion  82  of x-ray beam  16  is essentially unobstructed by patient  22  and impinges without beam collimation on measuring portions  60 . Data from one or both measuring portions  60  is used by imaging system  10  to compute a patient x-ray dose. In one embodiment, x-ray fan beam  16  has a thickness in a z-axis direction no greater than that of measuring portions or regions  60  to provide a measurement representative of the entire x-ray dose patient  22  receives. However, x-ray beam  16  has a greater thickness than collimated portion  80  to provide a full selection of image slices and thicknesses. Portion  82  of x-ray beam  16  is of greater thickness than post-patient beam collimated portion  80  impinging on imaging portion. Computer  36  receives measurements from measuring portion (or portions)  60  and utilizes these measurements to compute a radiation dose. In one embodiment, this computation is based on an assumption that dosage across the fan beam  16  thickness is uniform. A full thickness of portion  82  of beam  16  is used for the estimation. 
     Patient  22  is thus imaged utilizing radiation beam  16  and a post-patient collimated detector, i.e., imaging portion  50  of detector array  18 . Dose measurements are made utilizing a full thickness of unobstructed (i.e., non-beam collimated) portion  82  of radiation beam  16  and a detector having portions that are not beam collimated, i.e., measuring portions  60  of detector array  18 . 
     In one embodiment, there is no post-patient collimation, but detector array  18  has an imaging portion  50  that is not as thick as measuring portion or portions  60 . A thickness of radiation beam  16  that is as thick or thicker than imaging portion  50  of detector array  18  is measured by measuring portion or portions  60  to estimate patient radiation dose. 
     Although the invention has been described and illustrated in detail, it is to be clearly understood that the same is intended by way of illustration and example only and is not to be taken by way of limitation. In addition, the CT system described herein is a “third generation” system in which both the x-ray source and detector rotate with the gantry. Many other CT systems including “fourth generation” systems wherein the detector is a full-ring stationary detector and only the x-ray source rotates with the gantry, may be used if individual detector elements are corrected to provide substantially uniform responses to a given x-ray beam. Also, the present invention may be used with systems performing either axial or helical scans, or both types of scans. Accordingly, the spirit and scope of the invention are to be limited only by the terms of the appended claims.