IL109143A

IL109143A - X-ray detector for a low dosage scanning beam digital x-ray imaging system

Info

Publication number: IL109143A
Application number: IL10914394A
Authority: IL
Inventors: Wilent John William; Moorman Jack Wilson; Skillicorn Brian; Fiekowsky Peter Joseph
Original assignee: Cardiac Mariners Inc
Priority date: 1993-04-05
Filing date: 1994-03-28
Publication date: 1999-03-12
Also published as: JP2003265457A; JP4229859B2; IL109143A0; AU6700694A; CN1127553A; WO1994023458A3; JPH08508431A; JP2004261605A; CN1041237C; EP0693225A1; WO1994023458A2; JP3569526B2; EP0693225A4

Description

X-RAY DETECTOR FOR A LOW DOSAGE SCANNING BEAM DIGITAL "X"-RAY IMAGING SYSTEM THO ni o np'ioi* ·>κχ uta "x"->_np it>ma*o "x" np:) n»mn ιο*υκ> *τ> by THE APPLICANT:- CARDIAC MARINERS, INCORPORATED 120-B ALBRIGHT NAY, LOS GATOS, CA 95030 U.S.A.

THE INVENTORS: 1. JOHN WILLIAM ILENT, (DISEASED) LAST ADDRESS: 139 VICTORIA LANE, APTOS, CALIFORNIA 95003 U.S.A. 2. JACK WILSOK MOORMAN, (U.S. CITIZEN) 136 PINTA COURT, LOS GATOS CALIFORNIA 95032 U.S.A. 3. BRIAN SKILLCORN (U.S. CITIZEN) 898 THE DALLAS AVENUE SUNNYVALE, CA 94087 U.S.A. 4. PATER JOSEPH FIEKOKSKY (U.S. CITIZEN) 952 SPRINGE ROAD LOS ALTOS, CA 94024 U.S.A.

" CAM-001 This patent application is made in the names of inventors John William Wilent, deceased, Jack Wilson Moorman, Brian Skillicorn and Peter Joseph Fiekowsky, all assignors to Cardiac Mariners, Inc., a California corporation.

SEEQIEIGAIIQ X-RAY DETECTOR FOR A LOW DOSAGE SCANNING BEAM DIGITAL X-RAY IMAGING SYSTEM BACKGROUND OF THE INVENTION 1. Field Of The Invention The present invention relates to diagnostic x-ray imaging equipment. More particularly, the present Invention relates to a real time scanning beam digital x-ray imaging system having improved resolution and reduced x-ray emissions provided by incorporation of a multi-apertured collimatlon grid and a segmented x-ray detector array. 2. The Prior Art Real time x-ray imaging is increasingly being required by medical procedures as therapeutic technologies advance. For example, many electro¬ physiologic procedures in cardiology, peripheral vascular procedures, urological procedures, and orthopedic procedures rely on real time x-ray imaging.

Unfortunately, current clinical real. time x-ray equipment produces high ' levels of x-ray exposure to both patients and attending staff. The F.D.A. has reported anecdotal evidence of acute radiation sickness in patients, and concern among physicians of excessive occupational exposure. (Radiological Health Bulletin, Vol. XXVI, No. 8, August 1992).

' ' CAM-001 A number of real time x-ray imaging systems are known in the prior art.

These include fluoroscope-based systems where x-rays are projected into an object to be x-rayed and shadows caused by relatively x-ray opaque matter within the object are displayed on the fluoroscope located on the opposite side of the object from the x-ray source. Scanning x-ray tubes have been known in conjunction with the fluoroscopy art since at least the early 1950s. Moon, Amplifying .and Intensifying the Fluoroscopic Image by Means of a Scanning X-ray Tube. Science, October 6, 1950, pp. 389-395.

Scanning beam digital x-ray imaging systems are also well known in the art. In such systems έίη x-ray tube is employed to generate X-ray radiation. Within the x-ray tube, an electron beam is generated and focussed upon a small spot on the relatively large anode (transmission target) of the tube, inducing x-ray radiation emission from that spot. The electron beam is deflected (electromagnetically or electrostatically) in a raster scan pattern over the entire anode. detector is placed at a distance from the anode of the x-ray tube. The detector converts x-rays which strike it into an electrical signal in proportion to the detected x-ray flux. When an object is placed between the x-ray tube and the detector, x- rays are attenuated and scaUered by the object in proportion to the x-ray density of the object. While the x-ray tube is in the scanning mode, the signal from the detector is modulated in proportion to the x-ray density of the object.

Examples of prior art scanning beam digital x-ray systems include those described in United States Patent No. 3,949,229 to Albert; United States Patent - 2 - I ■ ' CAM-001 No. 4,032,787 to Albert; United States Patent No. 4,057,745 to Albert; United States Patent No. 4,144,457 to Albert; United States Patent No. 4,149,076 to Albert; United States Patent No..4, 196,351 to Albert; United States Patent No. 4,259,582 to Albert; United States Patent No. 4,259,583 to Albert; United States Patent No. 4,288,697 to Albert; United States Patent No. 4,321 ,473 to Albert; United States Patent No. 4,323,779 to Albert; United States Patent No. 4,465,540 to Albert; United States Patent No. 4,519,092 to Albert; and United States Patent No. 4,730,350 to Albert.

In a typical prior art embodiment of a scanning beam digital x-ray system, an output signal fronMhe detector is applied to the z-axis (luminance) input of a video monitor. This signal modulates the brightness of the viewing screen. The x and y inputs to the video monitor are derived from the same signal that effects deflection of the x-ray signal of the x-ray tube. Therefore, the luminance of a point on the viewing screen is inversely proportional to the absorption of x-r^ys passing from the source, through the object, to the detector.

Medical x-ray systems are operated at the lowest possible x-ray dosage level that is consistent with the resolution requirements for the procedure.

Accordingly, both dose and resolution are limited by the signal to noise ratio.

Time and area distributions of x-ray photons follow a Polsson distribution and have an associated randomness which is unavoidable. The randomness is expressed as the standard deviation of the mean flux, and equals its square root.

' CAM-001 The signal to noise ratio of an x-ray Image under these conditions Is therefore equal to the mean flux divided by the square root of the mean flux. I.e., for a mean flux of 100 photons, the noise is +/- 10 photons, and the signal to noise ratio Is 10.

Accordingly, the spatial resolution and the signal to noise ratio of x-ray images formed by scanning x-ray imaging systems are dependent, to a large extent, upon the size of the sensitive area of the detector. As the detector aperture is increased in area, more of the diverging rays are detected, effective sensitivity increases and the signal to noise ratio is improved. At the same time, however, the larger detector aperture reduces attainable spatial resolution as the "pixel" size (measured at the jplane of the object to be imaged) becomes larger. This is necessarily so because most objects to be imaged in medical applications (e.g., structures internal to the human body) are some distance from the x-ray source. In the prior art, therefore, the detector aperture size must be selected so as to effect a compromise between resolution and sensitivity, it not being possible to maximize both resolution and sensitivity simultaneously.

In medical imaging applications, patient dosage, frame rate (the number of times per second that the object is scanned and the image refreshed), and resolution of the image of the object are key parameters. A high x-ray flux may easily yield high resolution and a high frame rate, yet result in an unacceptably high x-ray dosage to the patient and attending staff. Similarly, low dosage may be achieved at the cost of an invisible image or an inadequate refresh rate. A successful medical imaging system must provide low dosage, high resolution and CAM-001 an adequate refresh rate of up to at least about 15 images per second - all at the same time. Therefore, systems such as the prior art scanning beam digital x-ray imaging system described above will not work with diagnostic medical procedures where exposure times are relatively long and where, as Is always the case with real patients, the x-ray dose received by the patient must be kept to a minimum.

It Is therefore an object of the present Invention to provide a scanning beam digital x-ray imaging system capable of use in medical diagnostic procedures undertaken on living human patients.

It is also fain object of the present invention to provide a scanning beam digital x-ray imaging system which provides high resolution images at adequate frame rates while minimally exposing the object under investigation to x-ray radiation.

It is a further object of the present invention to provide a scanning beam digital x-ray imaging system having improved resolution at a distance from the plane of the source of the x-rays while maintaining decreased x-ray flux levels.

BRIEF DESCRIPTION OF THE INVENTION A scanning beam digital x-ray imaging system ("SBDX") according to the present invention includes an x-ray tube having an electron beam source and a target anode. Circuitry is provided for focussing the beam and directing or scanning the beam across the target anode in a predetermined pattern. For - 5 - ί CAM-001 example, the predetermined pattern may be a raster scan pattern, a serpentine or *S" shaped pattern, a spiral pattern, a random pattern, a gaussian distribution pattern centered on a predetermined point of the target anode, or such other pattern as may be useful to the task at hand.

A collimating element, preferably in the form of a grid, may be interposed between the x-ray source and an object to be x-rayed. The collimating element may, for example, comprise a round metal plate having a diameter of about ten inches and including an array of apertures numbering 512 by 512 at the center row and column of the plate. As presently preferred, the grid may comprise a plurality of thin plates laminated together, each of the plates containing an array of apertures positioned to form an array of stepped apertures having desired orientations through the laminated structure. The collimating element Is preferably placed immediately in front of the emitting face of the x-ray tube. Other configurations of collimating element could also be used. Intone preferred embodiment of the present invention each of the apertures in the plate or grid is constructed so that it is directed toward (or points at) a detection point on a plane located a selected distance from the collimating element. That distance Is selected to allow placing the object to be x-rayed between the collimating element and the detection point. The function of the collimating element is to form thin beams of x- rays, arrayed like pixels, and all directed from a point on the anode of the x-ray tube toward the detection point.

A segmented detector array containing an x N matrix of detector elements ' CAM-001 (preferably a square array) is centered at the detection point. The detector array preferably comprises a plurality of densely packed x-ray detectors. Such an array can be designed, positioned and applied according to the present invention in a manner that yields high sensitivity without loss of resolution, resulting in an x-ray system having a resolution comparable to or better than that of prior art x-ray units at a dosage at least an order of magnitude less than that of prior art x-ray systems. This feature of the present invention has important implications in medical and other fields. Exposure to patients and attending medical staff involved in current procedures will be reduced. Procedures now impossible due to the radiation exposure risk will become possible. 1 The output of the detector array is an intensity value for each element of the detector array at each point in time that the x-ray beam is emitted through an aperture in the grid. Because each aperture is located at a different point in space relative to the object under investigation and the detector array^ a different output will be available from the detector array for each aperture that the x-ray beam travels through. The detector array outputs may be converted into an image in a number of ways. One method is to perform a simple convolution upon the array output, i.e., summing the intensity values of the array elements corresponding to each aperture scanned, and then normalizing. The output array could then be used to drive a video or other display. More preferred are the multi-image convolution and the multi-output convolution methods, described below, which provide enhanced visual output.

' ' CAM-001 The SBDX Imaging system of the present invention is also capable of stereo imaging where a collimation grid having two groups of apertures is used. In this case, one group of apertures is constructed to point to a first detection point where a first segmented detector is located and a second group of apertures is constructed to point to a second detection point where a second segmented detector is located. By constructing two images from the two segmented detectors and using conventional stereoscopic display methods, a stereo image may be produced.

The SBDX imaging system of the present invention is also capable of highlighted imagirig of materials which exhibit different x-ray transmissivities at different x-ray photon energies. Accordingly, for example, microcalcification, a precursor of breast cancer, may be imaged. By constructing the grid and/or anode to emit two or more groups of beams of x-rays each having different x-ray energy spectra, and directing each group to the detector array (plural defector arrays could also be used), the difference of transmissivities of the object under investigation at the various x-ray photon energies can be turned into an image, thus highlighting only those materials within the object under investigation which exhibit differential x-ray transmissivity. Optimized for the detection of calcium, for example, such an imaging system is a powerful tool for the early detection of breast cancer and other anomalies.

Utilizing a segmented array which intercepts the entire collimated x-ray beam and image processing the array output provides maximum sensitivity without ' · CA -001 sacrificing the resolution provided by using small surface area detectors. A non-segmented detector of the same size as the segmented array would provide the same sensitivity at lower resolution.

The system described herein is capable of being used in conjunction with the "Catheter Including An X-ray Sensitive Optical-Sensor Locating Device" described in U.S. patent application serial number 08/008,455 (CAM-003) filed January 25, 1993 and which is hereby incorporated herein by reference. U.S. patent application serial number 08/008,455 is owned by the assignee herein.

I BRIEF DESCRIPTION OF THE DRAWINGS Fig. 1 is a diagram showing the basic components of a low dosage scanning beam digital x-ray imaging system. .

Fig. 2 is a diagram showing the distribution of x-rays frorn an SBDX system in the absence of a collimation grid.

Fig. 3 is a magnified view of the grid and anode of an x-ray tube for a low dosage scanning beam digital x-ray Imaging system.

Fig. 4 is a diagram of an x-ray tube tor a low dosage scanning beam digital x-ray imaging system.

Fig. 5 is a cross sectional diagram showing the fabrication of an x-ray tube * CAM-001 or a low dosage scanning beam digital x-ray imaging system.

Fig. 6 is a diagram of a stereoscopic scanning beam digital x-ray imaging system.

Fig. 7A is a diagram of an apertured x-ray source interacting with a simple non-segmented detector.

Fig. 7B is a diagram of x-rays from a single aperture of an apertured x-ray source interacting with a segmented detector array. 1.

Fig. 7C is a diagram of x-rays from a number of apertures of an apertured x-ray source interacting with a segmented detector array.

Fig. 7D is a diagram of x-rays from two apertures of an x'-ray coliimation grid interacting with an object under investigation and then a segmented detector array.

Fig. 8 is a diagram of the exposed surface of a 5 x 5 detector array for a low dosage scanning beam digital x-ray imaging system.

Fig. 9 is a diagram of a 5 x 5 detector array for a low dosage scanning beam digital x-ray imaging system.

Fig. 10 is a diagram of a detector element for a low dosage scanning beam CAM-001 digital x-ray imaging system.

Fig. 1 1 is a diagram showing an array of pencil-type detector elements for a non-planar detector array.

Fig. 12 is a diagram of a 3 x 3 detector array for a low dosage scanning beam digital x-ray imaging system.

Fig. 13 is a diagram showing the basic components of a low dosage scanning beam digital x-ray imaging system utilizing negative feedback to control x-ray flux. i.

Fig. 14 is a perspective diagram showing the grid seal assembly.

DETAILED DESCRIPTION OF A PREFERRED EMBODIMENT Those of ordinary skill in the art will realize that the following description of the present invention is illustrative only and not in any way limiting. Other embodiments of the invention will readily suggest themselves to such skilled persons.

System Overview Turning to Fig. 1 , a scanning beam digital x-ray imaging system according to a preferred embodiment of the present invention is diagrammed. A scanning x-ray tube 10 is used as the x-ray source. An approximately -100kV to -120kV power - 1 1 - i ' ' CAM-001 supply is used to power x-ray tube 10 as is well known in the art. A 100 kV power supply will provide a spectrum of x-rays ranging to 100keV. As used herein, 100kV x-rays refers to this spectrum. X-ray tube 10 includes a deflection coil 20 under the control of scan generator 30 as is well known in the art. An electron beam 40 generated within x-ray tube 10 is scanned across a grounded anode 50 within x-ray tube 10 in a predetermined pattern. For example, the predetermined pattern may be a raster scan pattern, a serpentine or "S" shaped pattern, a spiral pattern, a random pattern, a gaussian distribution pattern centered on a predetermined point of the target anode, or such other pattern as may be useful to the task at hand. Presently preferred is the serpentine or "S" shaped pattern which eliminates the need in a raster sdan pattern for "fly back." As electron beam 40 strikes anode 50 at point 60, a cascade of x-rays 70 is emitted and travels outside of x-ray tube 10 toward the object 80 to be investigated with the x-ray. To optimize system performance, a cone of x-ray photons must be generated that will diverge in a manner that will just cover the detector array. This is preferably accomplished by placing a collimation grid between the anode of the scanning x-ray tube and the detector. Collimation grid 90 is therefore disposed between object 80 and x-ray tube 10. Collimation grid 90 is designed to. permit only those x-rays 100 which are directed toward detector 110 to pass through it. Thus, as electron beam 40 is scanned across anode 50, at any given moment, there can only be a single x-ray beam 100 passing from anode 50 to detector array 1 10.

* CAM-001 Fig. 2 depicts the distribution of x-rays in the absence of a coliimation grid.

The output of detector array 110 is then processed and may be displayed on monitor 120 as a luminance value at an x,y location on monitor 120 corresponding to the x,y location on anode 50. This may be accomplished by using the same scan generator to drive the x,y position of electron beam 40 and the position of the electron beam within the video monitor 120. Alternatively, image processing techniques can be used to produce a computer driven image on an appropriate display or photographic medium.

The X-Ray Tubeit Fig. 3 depicts a magnified view of the grid and anode structure. Anode 50 is preferably formed of tungsten deposited upon a beryllium anode support 130. Aluminum or other relatively x-ray transparent materials could be used as well. Tungsten is preferred for anode 50 because, having a relatively! high atomic number of 74, it readily emits x-rays when irradiated by an electron beam.

Tungsten's high melting point of 3370'C and good vacuum characteristics make it suitable for the high temperature and hard vacuum conditions within the tube. Beryllium is preferred for anode support 130 because it is strong and does not significantly attenuate or scatter the x-rays emitted from tungsten anode 50. The thickness of the tungsten layer which forms anode 50 is preferably 7 microns (the distance necessary to efficiently convert 100 kV electrons to x-rays). The thickness of beryllium anode support 130 is preferably about 5 mm. Anode support 130 should be as thin as possible subject to the physical constraint that it must be • CAM-001 strong enough to withstand the pressure gradient of one atmosphere across it.

Collimation grid 90 preferably consists of an array of apertures 140, each of which according to one preferred embodiment of the present invention Is oriented or pointed toward detector array 110. The number of apertures 140 in collimation grid 90 corresponds to the number of pixels, e.g., 512 by 512 to 1024 by 1024 at the center- of preferably round collimation grid 90 and, in part, determines the system resolution. The thickness of grid 90 and size of apertures 140 are determined by the distance of the detector array 1 10 from x-ray tube 10 (here, preferably 36 inches), the. need to attenuate all x-rays not heading to the detector, and the size of detector elements 160 of detector array 1 10. Apertures 140 are preferably laid out in a rectangular row and column pattern having a circular boundary ten inches in diameter. This is called the "circular active area". At the center of the circular active area the aperture count is preferably, according to one preferred embodiment of the present invention, 512 by 512. The, non-aperture portion 150 of collimation grid 90 is designed to absorb errant x-rays so that they do not illuminate object 80. This is accomplished by fabricating the grid so that x- rays striking the non-aperture portion 150 will see at least ten times the "1/2 value" (the amount of material necessary to attenuate 1/2 the x-rays striking it at the system energy, here, 100 keV). Errant x-rays would provide the object and attending staff with x-ray dosage but contribute no meaningful information to the image. Collimation grid 90 is preferably fabricated of 50 thin sheets of 0.010 inch thick molybdenum which are stacked and held together. Molybdenum is preferred because it readily absorbs x-rays so that x-rays generated by x-ray tube 10 which * ' " CAM-001 are not directed to detector 110 will be stopped before they impinge, uselessly and potentially harmfully, upon object 80, which, of course, may be a human patient. Lead or similar x-ray dense materials could also be used.

The apertures 140 of coliimation grid 90 are preferably square in cross section in order to obtain the maximum packing density and be compatible with the preferred square shape of the detector array elements 160. Other shapes could also be used, particularly hexagons. The square apertures 140 are preferably 0.015 inches by 0.015 inches in dimension which yields a cross sectional area that is about 1/100 the cross sectional area of conventional collimators used with fluoroscopes. Because of this tighter coliimation, a smaller beam width for x-ray beam 100 is achieved. This means that the cross sectional area of the face of the detector may be correspondingly much smaller than in conventional systems. As a result, x-rays scattered at the object miss the detector and do not fog the image as they do in conventional systems which utilize relatively large surface area detectors.

A preferred method for fabricating the coliimation grid 90 is by photochemical milling or etching. Photo-chemical milling Is presently preferred because it is cost effective and accurate. According to this method, a set of 50 photo masks is created to etch holes or interstices into 50 thin sheets of 0.010 inch thick material. The etched sheets are then stacked and aligned and held together to form a grid assembly having a plurality of stepped apertures, each of a predetermined angular relationship with respect to the sheets.

CAM-001 One presently preferred method for holding the etched sheets that grid 90 Is formed from is shown in Fig. 14. Etched sheets 91 (preferably 50) are each provided with alignment holes or alignment apertures 94. Alignment pegs 95 are placed in each alignment aperture 94 to align the etched sheets 91. The assembly of sheets 91 and pegs 95 is then placed in aluminum ring 359. Aluminum ring 359 is provided with a vacuum port 370 which may be sealed with pinch off 375.

Aluminum eheet 365 which is 1mm in thickness is then bonded and sealed with a vacuum adhesive to upper surface 380 of ring 359. Aluminum sheet 360 is similarly bonded to a lower surface 385 of ring 359. A partial vacuum is then pulled through port 370 and the port 370 is then sealed at pinch off 375 as is well known in the art. In this fanner, relatively x-ray transparent aluminum sheets 360, 365 provide a clamping action tending to hold etched sheets 91 together and in alignment as grid assembly 90.

The apertures 140 furthest from the center of grid 90 have 'a stepped surface and are preferably square in cross section. X-rays are generally unaffected by the roughness of the channels due to the stepped surface, and even if they are scattered, they will not measurably affect the resultant beam. The material used for the collimalion grid 90 as discussed above can be molybdenum, brass, lead, or copper with molybdenum presently preferred. Presently preferred tolerances for the positions of the holes are +/- 0.0005 inches center to center without cumulative error and for the hole sizes are +/- 0.001 inch.

Alternative methods for fabricating collimation grid 90 which could be used CAM-001 include electron beam machining, drilling or minimachining, and laser drilling, Drilling and laser drilling have the drawback that they generate round rather than square holes. Circular apertures are not presently preferred although they should work as well.

More details of the preferred scanning x-ray tube 10 are shown at Figs. 4 and 5. Electron gun 161 is located opposite the face of x-ray tube 10 and is operated at a potential of up to about -100kV to -120kV. Grounded anode 50 is located at the face of the tube and an electron beam 40 travels between electron gun 161 and anode 50. A grounded electron aperture plate 162 is located near electron gun 161 and ncludes an aperture 163 at its center for electron beam 40 to pass through. A magnetic focus lens 164 and deflection coil 20 position the beam spot on anode 50 utilizing dynamic focussing as is well known in the art. The tube is fabricated to have a 10 inch diameter circular active area in which electron beam 40 may intersect anode 50 with the electron beam deflected up to ab'put 30' at the extremities of the circular active area. When the beam is not being "fired" through a particular aperture it is preferably left off, resulting in a power savings of up to about 25%.

Turning to Fig. 5, a cross sectional view of the front portion of x-ray tube 10 is depicted. The interior of the x-ray tube 340, maintained at a vacuum is rearward of anode 50. Anode 50 is a 7 micron thick coating of tungsten. Forward of anode 50 is beryllium anode support 130 which is 5 mm thick. Forward of beryllium anode support 130 is cooling jacket 350 which is preferably 4 mm thick and may be ' CAM-001 adapted to carry water or forced air. Aluminum grid supports 360, 365 are each 1 mm thick and help support collimation grid 90 which is preferably 0.5 inches thick.

Stereoscopic X-ray Imaging Turning now to Fig. 6, according to another preferred embodiment of the present invention, a grid having more than one focal point may be provided so that stereoscopic x-ray images may be obtained. If, for example, every other row of apertures in grid 90 were pointed at focal point F1 (92) and the remaining apertures were pointed at focal point F2 (93), by placing a first sensor array at F1 (92) and a second sensor array at F2 (93) it is possible to scan the apertures in a raster or serpentine pattern and thereby create a "line" of data for the first sensor array, then a line of data for the second sensor array. Repeating this, it is possible to build up two complete images, as seen from two distinct points in space, F1 and F2, and thereby display them with conventional stereoscopic imaging display systems to provide a stereoscopic x-ray image. ·'( The Array Detector To achieve resolutions of several lines per millimeter at the object plane, as are required in some medical applications, the spatial resolution limit is in large part determined by the size of the detector. This is because, with today's x-ray tube technology, it is not feasible either to produce the very high power levels that would be required in order to obtain sufficiently intense highly directionalized x-ray emissions or to develop the associated x-ray directionalizing means.

' ' ' CAM-001 When the detector is made smaller than the area intersecting the cone of emitted x-rays, a large proportion of the x-rays emitted by source 50 miss detector 250, as shown in Fig. 7A. This is, in fact, how industrial scanning beam digital x-ray inspection systems are designed, where dose is usually not an issue. As a consequence, the dose is increased in order to maintain the desired resolution.

Accordingly, resolution improves with the use of smaller detectors, but x-ray dose is minimized when the area of the detector equals or exceeds the area defined by the cone of emitted x-rays intersecting detector plane 270.

Resolution pi a scanning x-ray imaging system is determined by the cross sectional area of a detector element projected onto the object plane 280 (the plane perpendicular to a line between the center of anode 50 and the center of detector 110 in which object 80 is located). Thus, if a large area detector is subdivided into smaller array elements, e.g., as shown in Fig. 8, the large capture^rea of the aggregate detector is maintained, while simultaneously retaining an image resolution that is proportional to the size of an individual small detector element 160.

The resolution defined by the individual detector elements 160 is maintained by distributing and summing the readings from the individual elements into a memory buffer in which each address, i.e., pixel, corresponds to a specific location in the object plane 280. As the x-ray beam 100 is moved discretely across collimating grid 90, which is positioned in front of the x-ray emitting anode 50, the ' ' CAM-001 address, to which the output of a given detector element is added, changes. The imaging geometry is shown in Fig. 7B and 7C. In Fig. 7B a single beam position is shown along with how it is divided among 5 pixels. In Fig. 7C the sequential beam positions are shown along with how they are added together within a single pixel.

In other words, the signal for each of the detector elements is stored in an image buffer, at a memory address that corresponds to a very small specific region in the object plane 280, i.e., a single pixel. Accordingly, the memory storage address for each detector element changes with the location of the scanning x-ray beam in an ordered fashion such that each pixel in memory contains the sum of the radiation passing through a specific spot in the object plane 280. In this way the resolution of the system is determined by the size of a single detector element, while the sensitivity of the system is optimized, since virtually all of the x-rays reaching the detector plane 270 are recorded. Ί An additional benefit of this array detector imaging geometry is that the object plane 280 is narrowly defined. Structures lying in front of or behind it will be blurred (out of focus). X-rays from a first aperture 141 and a second aperture 2 are depicted in Fig. 7D passing through an object plane 280 a distance SO from apertures 141, 142 and passing through a plane 281 twice distance SO from apertures 141, 142. As can be readily seen, the resolution degrades to about 1/2 that available at SO at the distance twice SO. This feature provides for improved localization and visualization of detailed structures in the plane of interest 280, while providing an adequate depth of field that may be modified by the system CAM-001 geometry.

Tha X-ray Detectors Conventional image intensifier technology has basic constraints that limit a system's sensitivity. The thickness at which the scintillator material can be applied is limited by Its optical transmission properties. Typically, it is made thick enough to capture about 50 percent of the incident x-ray photons. Of the emitted light photons, only about half reach the photocathode. At the photocathode, only about 10 percent of the incident light photons produce photoelectrons. Thus, only about 2.5 percent (.5 x .5 K .1) of the incident x-ray photon energy is conserved in an t- image intensifier system. In addition to this limited conversion efficiency, light photons are scattered laterally by the scintillator material and create haze that reduces the system's resolving power at a given dose level.

V One of the primary objects of the present invention is to provide an SBDX imaging system which will ensure that the subject under examination is exposed to the lowest possible level of x-rays commensurate with achieving image quality adequate to meet the requirements of the procedure being performed. This means that the system used to detect the x-ray photons emerging from the subject must have the highest possible photon to electrical signal conversion efficiency. In order to achieve this, the material used for the detector must have a length in the direction in which the photons travel that is sufficient to ensure that no photons emerge from the end farthest away from the incident x-rays, i.e., the photon energy ' CAM-001 must be adequately dissipated in the material in order to maximize the output of the detector. There are several types of detectors which could be used in the presently described SBDX system. That which is currently preferred is the scintillator i which x-ray photon energy is converted to visible light energy and the light intensity is then converted to an electrical signal by means of a photomultiplier, photo diode, CCD or similar device. Because each pixel in the SBDX image must be generated in a very short time period, about 140 nanoseconds, the scintillator material must have a fast response and a minimum afterglow time. Afterglow is the phenomenon wherein the scintillator continues to emit light after the stimulating incident x-radiation has ceased. Plastic scintillators, such as organic loaded polystyrene, are suitable in that they) have the required fast response characteristics but they have a relatively small x-ray photon interaction cross section so that their linear x-ray absorption coefficients are also small in value. The consequence is that a considerable thickness is required to stop ail the x-ray photons. For TOOkV x-rays, as presently preferred, a typical plastic scintillator must be about 28p mm (11 inches) thick to capture 99% of the incident x-rays.

According to a presently preferred embodiment of the present invention, the SBDX array detector 110 comprises a 9 by 9 array of 81 densely packed detectors 160 spaced at a distance of 914 mm (36 inches) from the x-ray source 50. (A 5 by 5 and a 3 by 3 array are also contemplated as is a non-square array having square detectors filling a circle about the x-ray target). The geometry of the collimator aspect ratio, its spacing from x-ray source 50 and the size of x-ray source 50 gives a square section pyramid of x-rays with a 1.6 degree total included angle which is ' CAM-001 about 25.7 mm (1.01 inches) across at the entrance to detector array 110. Each individual scintillator 170 must therefore be 2.85 mm across at its entrance face 172. If scintillator 170 has parallel sides, x-rays entering near to the edges will not be able to travel the requisite distance without striking the scintillator walls. These x-rays may therefore pass through to the adjacent scintillator if it is not shielded causing it to generate an output seemingly from the wrong spatial position in the subject with consequent degradation of the image quality. To avoid this effect, the individual scintillators are preferably tapered so that their bounding faces have the same included angle as that of the incident x-rays. In the preferred example cited above, each scintillator 170 would thus preferably be a frustrum of a pyramid 280 mm long, with an entrance face (172) 2.85 mm across and a face at the photodetector end 174 which is 3.7 mm across. The whole bundle of 81 detectors therefore has multifaceted ends, each facet being tangential to the surface of a sphere centered at x-ray source 50.

I A further improvement in detection efficiency for scintillators may be achieved by tapering the scintillator at an angle that Is greater than that of incident x-ray beam 100. Photoelectrons and scattered x-rays produced by interaction between the incident x-ray and scintillator atoms near to the edges of the scintillator may be lost to the shielding material which preferably separates adjacent scintillators. These tost photoelectrons will not produce any light so they will not contribute to the light output from the scintillator. Their loss therefore reduces the efficiency of the scintillator. The maximum distance travelled by a photoelectron depends upon its energy and on the material in which it is travelling. For 100kV x- CAM-001 rays interacting with the atoms in a plastic scintillator, there will be no photoelectrons which can travel a distance greater than about 0.1 mm. If the scintillator pyramid frustrum is made with an included angle greater than that of x-ray beam 100 so that its dimensions become greater than the beam envelope by 2 x 0.1 mm, in a distance short compared with the detector length (280 mm) then the efficiency reduction due to the lost photoelectrons will be minimal. In this case, the center of the sphere tangential to the facets would no longer coincide with x-ray source 50 but would be nearer to detector array 110.

Scattered photons will travel greater distances than the photoelectrons; so, to prevent these frftrn escaping to an adjacent scintillator the rate of taper of the scintillator pyramid may be made greater than is needed for complete photoelectron capture in order to maximize the scattered photon capture.

Referring now to Fig. 9, according to a presently preferred embodiment of the present invention, in contact with each scintillator element 170 is a light pipe or fiber optic cable 180 which optically couples each scintillator element 170 with a corresponding photomultiplier tube 190 or solid state detector. Alternatively scintillators 170 may be located in close physical proximity to appropriate photodetectors.

Fig. 10 shows a preferred configuration of a detector element 160. An x-ray opaque sheet 200 with apertures 210 corresponding to each detector element 160 is disposed in front of detector array 110. Each detector element 160 is enclosed In CAM-001 a light tight enclosure 220 which may also be x-ray opaque. A light blocking window 230, preferably made of thin aluminum sheet is located at the front of light tight enclosure 220. Light blocking window 230 is x-ray transparent. Within light tight enclosure 220 is a scintillator element 170 in close proximity to a photomulliplier tube 190 which is electrically connected to a pre-amplifler 240. Preferably the analog signal from the pre-amplifier 240 is converted to a digital signal in a conventional manner for further processing.

Alternatively, scintillators could be placed in direct or close contact with an array of photo diodes, photo transistors or charge coupled devices (CCDs) to achieve a more bugged and compact detector. Where solid state devices, particularly CCDs, are used, cooling, such as with a Peltier-type cooler, or the like, may be employed to increase the signal to noise ratio of the device.

Alternatively, the scintillator array could be placed in direqt or close contact with one or more position sensitive photomulliplier tubes which provide an output signal which identifies the position coordinates of the light source as well as its amplitude.

In another preferred embodiment, the sensor array may comprise a collection of pencil-type detectors 285 arrayed as shown, for example, in Fig. 11. In Fig. 11 tapered scintillators 290 are arrayed in the path of x-ray beam 100 so that the scintillator corresponding to a particular cross-sectional area of beam 100 will fully absorb x-rays within that cross-sectional area. Photo multiplier tubes 300 are ' ' CA -001 located in close physical proximity to scintillators 290 so that an electrical signal will be generated in response to the absorption of x-rays by scintillators 290. Solid state devices could also be used In place of photo multiplier tubes 300.

According to a presently preferred embodiment of the present invention, the long sides of the plastic scintillators are coated with a reflective material such as aluminum to prevent light from escaping (or entering) and to aid in total internal reflection within the scintillators.

According to another preferred embodiment of the present invention, each scintillator element. 179 is isolated from its adjacent scintillator elements 170 by a thin sheet 171 of a highly x-ray opaque material such as, for example, gold or lead. Sheets 171 may preferably be about 0.004" to 0.005" thick. The position of sheets 171 between the scintillators 170 is shown in Fig. 12.

Image Processing An important refinement of the present invention concerns the application of an image processing system to further reduce the required dosage. In practice, the signal from the detector is not usually applied directly to the "z" or luminance input of a video monitor. Instead, digitized intensity data for each pixel are stored in a discrete address in a "frame store buffer". More than one such buffer may be used in certain applications. Pixel addresses within the buffer can be randoml accessed and the numeric intensity value can be manipulated mathematically. This function has application in applying various image enhancement algorithms CA -001 and it allows for pixel assignment of the data from discrete segments of the detector array.

According to a preferred embodiment of the present invention an SBDX image would consist of up to about 262,144 pixels, arranged in 512 rows and 512 columns (corresponding to the 512 rows by 512 columns of apertures at the center of collimation grid 90). For the purpose of the explanatory example below, it is assumed that the scanning x-ray source Is momentarily centered upon the pixel, P, located at row 100 and column 100. It is further assumed that the detector array consists of a 3 by 3 array containing 9 segments (Fig. 12) and that each segment is sized so as to intercept all of the x-ray emissions associated with a single pixel.

The numerical values, digitized from the individual segments of the detector array, are assigned to pixel addresses as follows: Segment 1 -- row 99, column 99 '·', Segment 2 -- row 99, column 100 Segment 3 -- row 99, column 101 Segment 4 ~ row 100, column 99 Segment P row 100, column 100 Segment 6 -- row 100, column 101 Segment 7 - row 101, column 99 Segment 8 - row 101, column 100 Segment 9 -- row 101, column 101 The same pattern of data assignment is repeated as the scanning x-ray beam ' ' CAM-001 passes all of the pixels.

In the displayed image, the numerical value of each pixel is equal to the sum of "n" parts where "n" is the number of segments in the array (in this example, n=9).

When constructed as shown herein, the detector array has the effect of fixing the working distance at which optimum focus is obtained and providing a plane of optimum focus not available in prior art non-segmented detector array SBDX imaging systems. The following parameters must be taken into consideration in design of the detector: 1. Ίριβ size and shape of the collimated beam from the x-ray source 50; 2. The distance between the source 50 and the detector array 110, "SD"; 3. The distance between the source 50 and the center of the object of interest 80, "SO"; 4. The desired resolution, or pixel size at the object of interest 80; 5. In medical applications, the total area of the array must be large enough to intercept all of the x-rays from the collimated focussing grid 90.

In an SBDX system according to a presently preferred embodiment of the present invention, the distance between the x-ray source 50 and the exit side 260 of focussing grid 90 is about 0.894 inches (see Figs. 3, 5). Apertures 140 are 0.015 inches by 0.015 inches square. The spot size of electron beam 40 on anode 50 is » · CAM-001 about 0.010 inches in diameter. The detector array is 36 inches from anode 50. Thus, the beam width of x-ray beam 100 is 2*ARCTAN((spot diameter/2)/((aperture width/2)+(spot diameter/2))*0.894 inches, or 1.6*. At a distance of 36 inches from anode 50, the projected x-ray beam diameter is 36*TAN(1.6*). Therefore, the detector array 110 should be at least about 1 inch on a side for this preferred embodiment. For example, if the object to be imaged is 9 inches from anode 50 and the desired pixel size is 0.020 at the object, and the distance from source to detector, again, is 36 inches with an optimal detector array size of 1 inch square, the projected size of pixels at the detector plane 270 is simply (SD/SO)*pixel size at object, or 0.080 inches. Dividing 1 inch by 0.080 inches we see that the desired resolution may be bbtained with a square segmented detector having 12 to 13 segments on a side. Obviously, many other configurations could be used depending upon the circumstances in which the SBDX system is to be used.

Outside of the plane of optimum resolution, SO (280 in Fig ,7D), resolution will degrade to one half at 0.5 x SO and at 2 x SO (281 in Fig. 7D). This allows for a reasonable depth of focus for most applications. In some applications, such as imaging the human heart, degraded focus outside of this range of depth is seen as being advantageous. Blurring of detail outside of the area of interest tends to increase the perception of details within the area of interest.

A number of methods can be used to obtain a useable image from the data obtained as described above. As described above, a simple convolution may be used, however, in this case, resolution will not be fully optimized. Two additional ' CAM-001 methods are presently preferred for obtaining maximal resolution and sensitivity from the captured data. These are called the multi-image convolution method and the multi-output convolution method. For both cases, the following is assumed: There are NCQ columns and NRG rows in grid 90. Each intersection of a column and row is a "pixel." Those pixels outside of the circular active area of grid 90 are treated as if they contribute no measured intensity to the image. Pixels not illuminated by x-ray beam 100 during a scan are similarly treated as if they contribute no measured intensity to the image. There are N columns of sensor elements 160 and M rows of sensor elements 160 in a rectangular or square sensor array 110 of lcjimensions M by N. ZRATIO is a real number between 0 and 1. If ZRATIO=1, the focus is set at the sensor plane. If ZRATIO=0, the focus is set at the x-ray source plane. If ZRATIO=0.5, the focus is half way between the x-ray source plane and the sensor plane, and so on. PIXELRATIO is the number of image pixels per physical distance between adjacent sensors in a column or row. For example, if the spacing between pixel centers at objec plane 280 is 1.0 mm, and the spacing between sensors at detector plane 270 is 10.0 mm, then PIXELRATIO=10. FOCUS=ZRATIO*PIXELRATIO. IMAGE is a data array of dimension M x N containing the intensity information for a particular scan and corresponding to a particular pixel. PIXEL is a 4 dimensional array of dimension NRG x NCG x M x N which contains the NRG x NCG IMAGE data arrays obtained by scanning all (or part of) the apertures. PIXEL is refreshed after each scan according to one preferred embodiment of the present invention.

CAM-001 As the beam is scanned across the anode surface, it is, in effect, positioned before the center of selected apertures 140, "fired," and then repositioned. Thus for each firing, an IMAGE array of data will be acquired. While these images could be constructed into a displayable image having some use directly, more resolution and sensitivity is obtained by combining them. The first preferred method for combining the images is called the multi-image convolution method. In the multi-image convolution method, an OUTIMAGE array of intensities of dimensions NRG x NCG, which can be displayed on a CRT or like display means, is formed by assigning to OUTIMAGE(y.x) the value of: j=N i=?M The second presently preferred method for combining the Nf jG x NCG IMAGE data arrays into a useful picture is called the multi-output convolution method. In this case, with a sensor array of M x N sensors there will be M x N digitizers (or their equivalents, multiplexed) and the same number of pixel summing circuits. The digitized values from each sensor are called SENSOR(|,i). The final OUTIMAGE array is computed as follows - for each pixel in the output Image array OUTIMAGE(y,x) [for y=1 to NRG and x=1 to NCG] one pixel from each of the M x N source images SENSOR(j.i) Is summed [for J=1 to N and 1=1 to M] into destination image pixel OUTIMAGE(y-]*FOCUS,x-i*FOCUS). Normalization Is then carried out over the OUTIMAGE array by dividing each element thereof by M*N. ' 1 CAM-001 A further improvement upon these techniques may be obtained by performing linear interpolation based upon the fractional part of the FOCUS factor.

An advantage of the multi-image convolution method over the multi-output convolution method is that the former allows the plane of optimum focus to be selected in software after the data is captured while the latter does not. The latter method, owever, may be performed more quickly where timing is a limitation.

Negative Feedback X-Ray Flux Control Turning now to Fig. 13 an SBDX imaging system employing a negative feedback path 305 to control the x-ray flux of x-ray beam 100 is depicted.

Preferably negative feedback from the sensor array is utilized to control x-ray flux so that the sensor array always sees approximately the same flux level. In this way when soft tissue (which is relatively transparent to x-rays) is being scanned, the x- ray flux will drop, reducing the overall dosage to the patient (or,object). Improved contrast and dynamic range are provided by using negative feedback flux control. According to this embodiment, differential amplifier 310 has an adjustable reference level 320 which may be set by the user. Negative feedback loop 305 feeds back to x-ray tube 10 to control the x-ray flux.

Time Domain Scanning Mode A time domain x-ray imaging system may also be implemented using the principles disclosed herein. In such a system, the time to reach a predetermined measured x-ray flux from the various pixels could be computed and mapped.

" ' CAM-001 Negative feedback control could then be employed to turn off or reduce x-ray flux from apertures corresponding to pixels which had reached the predetermined flux level for the scan period in question. In this case, the information gathered would be time to flux level and the mapped or imaged information would correspond to time rather than intensity. Such a system has the potential to provide much higher signal to noise ratios, improved contrast, drastically reduced x-ray dosage to the object under investigation, and improved dynamic range.

Multiple Energy Xrray Imaging Mode According to one preferred embodiment of the present invention, two or more groups of xi>ray beams 100 are directed toward one or more detector arrays. A first group of x-ray beams has a first characteristic x-ray energy spectrum. A second group of x-ray beams has a different second characteristic x-ray energy spectrum. By comparing the measured transmissivities of the first and second group of x-ray beams, the presence of certain materials in the object under investigation may be detected. The basic concept of use of differential x-ray imaging is known in the art and is disclosed, for example, in U.S. Patent No. 5,185,773 entitled "Method and Apparatus for Nondestructive Selective Determination of a Metal" which is hereby incorporated herein by reference.

The two groups of x-rays may be generated in a number of ways. One such way is by fabrication of a special anode 50 having a first material or first thickness of a material adjacent to the apertures of the first group of apertures and a second material or second thickness of material adjacent to the apertures of the second CAM-001 group of apertures. In this manner, the apertures associated with the first group will emit x-rays having a first characteristic energy spectrum and the apertures associated with the second group will emit x-rays having a second characteristic energy spectrum. Alternatively, K-filtering (or K-edge filtering) can be used by placing filter material (such as, for exampje, molybdenum) within a portion of the apertures 140 to produce a similar effect. In this case, a first group of apertures would comprise a first filter inserted therein and a second group of apertures would comprise a second filter inserted therein. The second filter could be no filter at all. As in the previous case, two groups of x-rays having different characteristic energy spectra would be associated with the two groups of apertures.

X f Once at least two groups of apertures are associated with different characteristic x-ray spectra, it is now possible to detect micro-calcification (a precursor of breast cancer) and other abnormalities not normally visible with broadband x-rays. For example, by performing a scan of the firsf group of apertures to form a first image, then performing a scan of the second group of apertures to form a second image, and subtracting the images to highlight their differences, it is possible to detect micro-calcification and other such abnormalities with a low dosage scanning beam x-ray imaging system - in real time. · Similarly, a multiple detector array arrangement could be used with group 1 apertures directed toward a first detector array and group 2 apertures directed toward a second detector array, etc.

While a number of preferred embodiments have been discussed above for CA -001 various configurations of the present invention, the following specifications are illustrative of a presently preferred SBDX imaging system according to the present invention Grid Shape: round Diameter: 10 inches Aperture Pitch: 0.020 inches Numbe of Apertures across a diameter: 500 Area of grid: 78.5 sq. inches Number of apertures: 196,350 approx.

Aperture cross-sectional shape: square Aperture width: 0.015 inches Space between apertures: 0.005 inches B. Source-Detector Distance: 36 inches Location of Pla¾e of Optimum Focus: 9 inches from Grid Scan Frequency: Adjustable to 30 Hz Operating voltage on x-ray tube: 70-1 OOkV Detector Array: Shape of detector elements: square Size of detector elements: 0.06 inches x 0.06 inches Number of detector elements: 49 in 7 x 7 array Total included angle subtended from the detector center point to the collimation grid outside diameter: 15.8° Field of view at Plane of Optimum Focus: 7.5 inches Pixel size at Plane of Optimum Focus: 0.38mm Pixel size at Detector Plane: 1.52mm Resolution: 1.3 linepairs/mm Accordingly, an SBDX imaging system utilizing a segmented detector array has been shown and described which simultaneously provides high resolution, high sensitivity, and low x-ray dosage to the object under investigation. The system also permits the point of optimum focus to be set at any point between the source CAM-001 50 and the detector array 110, and provides an effective working depth of field.

While embodiments and applications of this invention have been shown and described, it would be apparent to those skilled in the art that many more modifications than mentioned above are, possible without departing from the inventive concepts herein. The invention, therefore, is not to be restricted except in the spirit of the appended claims.

Claims

103, 143/2 C L A I M S ΐ

1. An x-ray imaging system comprising an x-ray source, an x-ray detector array, a monitor and a scan generator? said x-ray source comprising a charged particle beam generator, an anode and a collimator? said anode which emits x-ray photons when stimulated by a beam of charged particles; said scan generator comprising circuitry for controlling said charged particle beam generator; said collimator disposed between said anode and said x-ray detector array in close proximity to said anode; said x-ray detector array comprising a plurality of detector elements, said detector elements arranged in an a,tt®9 forming a generally planar detector surface having detector array comprising a plurality of /rows and columns; said collimator constructed of an x-ray absorbing material, said x-ray absorbing material comprising a plurality of apertures, said axis of each of said apertures intersects said detector surface substantially at said center of said x-ray detector array and intersects said target layer; said anode, said collimator and said x-ray detector array positioned such that when said charged particle beam stimulates said target layer at a first point of intersection of a first one of said axis of said apertures and said target layer, said emitted x-ray photons pass through said first aperture and strike said detector elements of said x-ray detector array and when said charged particle beam stimulates said target layer at a second point of intersection of a second one of said axis of said apertures and said target layer, said emitted x-ray photons pass through said second aperture and strike all of said detector elements of said 109,143/2 x-ray detector array, said scanning beam generator successively positions said charged particle beam at said intersections of said axis and said target layer in a predetermined pattern; each of said detector elements comprising means for converting the energy from said x-ray photons which strike said detector elements to electrical signals, said electrical signals from each of said detector elements are discretely output from said x-ray detector array to said monitor*, a group of discrete electrical outputs is output for each positioning of the charged particle beam at successive intersections of said axes and said target layer, said monitor comprising circuitry which processes sa d groups of discrete signals to generate a visual image based upon said groups of discrete electrical signals.

2. An x-ray imaging system as claimed in claim l, wherein said detector elements each comprise a scintillator coupled to an input end of an optical fiber, an output end of said optical fiber is coupled to a means for converting the energy from said light photons to said electrical signals.

3. An x-ray imaging system as claimed in claim 1, wherein said detector elements are disposed in an array approximating a shape of a circle.

4. An x-ray imaging system as claimed in Claim 2, wherein said scintillator comprises a material selected from the group consisting ofi YSO, LSO, BOO and plastic. o 109,143/2

5. An x-ray imaging system as claimed in Claim 2, wherein said scintillator has a generally rectangular shape .

6. ¾. An x-ray imaging system as claimed in Claim l, wherein said detector array is substantially square.

7. An x-ray imaging system as claimed in Claim 1, wherein a first face of each of said detectors faces said s-ray source and a second face of each of said detectors on a portion of said detector opposite said first face, so that said first face comprises less area than said second face.

8. An x-ray imaging system as claimed in claim 1, wherein a portion of each of said scintillators is coated with a reflective coating.

9. An x-ray imaging system as claimed in Claim 1, further comprising means for creating a multi-output convolution image based upon said electrical signals, coupled between said detector array and said monitor.

10. An x-ray imaging system as claimed in Claim 1, further comprising means for creating a multi-image convolution image based upon said electrical signals, coupled between said detector array and said monitor.

11. An x-ray imagin system as claimed in Claim 1, further comprising means for creating a convolution image based upon said electrical signals, coupled between said detector array and said monitor. 109, 143/2

12. An x-ray imaging system as claimed in Claim 1, wherein said plurality of apertures comprises a first plurality of apertures and a second plurality of apertures, and wherein said x-ray detector array comprises a first detector array and a second detector array each having a center, so that said first array of apertures each has an axis that intersects said center of said second array and said second array of apertures each has an axis that intersects said center of said second array of apertures.

13. An x-ray imaging system substantially as described in the specification and in any one of Claims 1 to 12. For the Applicants, Simon Lavie Patent Attorney