Radiographic system

Abstract: A radiographic system is described which includes an X-­ ray source, and an x-ray detector including an x-ray sensitive surface. The x-ray sensitive surface is curved so as to eliminate differences in the lengths of the x-ray path between the source and different points on the surface.

Description: This invention relates generally to the field of radiographic systems. Radiography is a long known medical diagnostic imaging technique. In a conventional radiography system, an x-ray source is actuated to direct a divergent area beam of x-rays through a patient. A cassette containing an x-ray sensitive phosphor screen and light and x-ray sensitive film is positioned in the x-ray beam on the side of the patient opposite the source. X-radiation passing through the patient's body is thereby attenuated in various degrees to produce on the film a shadow image of a portion of the patient through which the x-rays pass. More recently, digital radiographic techniques and systems have been developed. In digital radiography the source directs x-radiation through a patient's body to a detector assembly located in the beam path beyond the patient. The detector produces electrical signals defining the radiation pattern emergent from the patient and incident on the assembly. These signals are then processed to yield a visual display of the image. The detector assembly includes an elongated planar array of individual detector elements. A detector element can suitably comprise a scintillator having a receiving face positioned in front of a photodiode. Each detector element responds to incident x-radiation to produce an analog electrical charge signal indicative of such radiation. These analog electrical signals represent the radiation pattern emergent from the patient's body. The analog signals are sampled and processed by imaging circuitry, primarily to improve their signal to noise ratio, and are subsequently digitized. The digital signals are fed to a digital data processing unit. The data processing unit records and/or processes and enhances the digital data. A display unit responds to appropriate digital data representing the image to convert the digital information back into analog form and to produce a visual display of the patient's internal body structure derived from the acquired image pattern of radiation. The display unit can be coupled directly to the digital data processing unit for substantially real time imaging, or can be fed stored digital data from digital storage means such as tapes or disks, representing patient images produced from earlier studies. Digital radiography includes techniques in which a thin spread beam of x-radiation is used. In practice of this technique, often called "scan (or slit) projection radiography" (SPR), the spread beam is scanned across the patient, or the patient is movably interposed between the spread beam x-ray source and the detector assembly, the detector being maintained in continuous alignment with the beam. The relative movement effected between the source-detector arrangement and the patient's body scans a large portion of the body. Discrete element detectors have been proposed comprising a single line of detector elements. Other proposals have included planar rectangular detector arrays of square detector elements. Details of certain aspects of digital radiography systems such as described here are set forth in the following publications: Mattson, R.A., et al, "Design and Physical Characteristics of a Digital Chest Unit", S.P.I.E. Volume 314, Digital Radiography (1981). Arnold, B.A. et al "digital Radiography: An Over-view" Proceedings of S.P.I.E. Volume 273, March 1981; Kruger, R.A. et al "A Digital Video Image Processor for Real Time X-ray Subtraction Imaging" Optical Engineering, Volume 17, No. 6 (1978); United States Patent No. 4,383,327, issued on May 10,1983, to Kruger. European Patent Application Publication No. EP 0115125-Al, published August 8, 1984, by Gary L. Barnes and entitled "Split Energy Level Radiation Detection"; United States Patent Application Serial No. 542,384, filed October 17, 1983 by Mattson, R.A.,et at entitled "Improving Signal Characteristics in Digital Scan Projection Radiography", owned by the assignee of this application, and corresponding to patent family member EP-A-0 138625 published April 24, 1985. It has also been proposed, where the detector array comprises a planar rectangular array of square detector elements, to improve the signal to noise ratio of the information developed by the detector, by the use of time delay and integrate (TDI) circuitry. An embodiment of such a proposed system is described in United States Letters Patent No. 4,383,327, issued on May 10, 1983 to Kruger. Such proposed TDI systems employ sampling at regular intervals of detector motion, and motion-synchronous shifting and adding of the individual detector-produced analog charge signals which are sampled. An important technique for enhancing a digitally represented image is called "subtraction". There are two types of subtraction techniques, one being "temporal" subtraction, the other being "energy" subtraction. Temporal, sometimes called "mask mode" subtraction, is a technique that can be used to remove overlying and underlying structures from an image when the object of interest is enhanced by a radiopaque contrast agent. Images are acquired with and without the contrast agent present and the data representing the former image is subtracted from the data representing the latter, substantially cancelling out all but the blood vessels or anatomical regions containing the contrast agent. A principal limitation of digital temporal subtraction is the susceptibility to misregistration, or "motion artifacts" caused by patient movement between the acquisition of the images with and without the contrast agent. An alternative to temporal subtraction, which is less susceptible to motion artifacts, is energy subtraction. Whereas temporal subtraction depends on' changes in the contrast distribution with time, energy subtraction exploits energy-related differences in attenuation properties of various types of tissues, such as the difference of the attenuation characteristics of soft tissue and bone. Soft tissue shows less change in attenuation capability with respect to energy than does bone. This phenomenon enables performance of energy subtraction. In practising that technique, pulses of x-rays having alternating higher and lower energy levels are directed through the patient's body. When a lower energy pulse is so generated, the detector and associated digital processing unit cooperate to acquire and store a set of digital data representing the image produced in response to the lower energy pulse. A very short time later, when the higher energy pulse is produced, the detector and digital processing unit again similarly cooperate to acquire and store a separate set of digital information representing the image produced by the higher energy pulse. The values obtained representing the lower and higher energy images are then processed in accordance with techniques described in the following publication: Lehmann, L.A. et al, "Generalized Image Combination in Dual KVP Digital Radiography" Medical Physics Volume 8, pp. 659-667 (1981). By processing in this manner, the image contrast and visibility of different tissues is substantially enhanced. Energy subtraction has the advantage, relative to temporal subtraction, of being substantially not subject to motion artifacts resulting from the patient's movement between exposures. The time separating the lower and higher imaging acquisitions is quite short, often less than one sixtieth of a second. An important disadvantage in dual energy subtraction techniques results from the necessity of rapidly alternating the output of an x-ray tube between high and low levels. This requirement gives rise to severe problems in a practical clinical device. The switching frequency is required to be on the order of 500 Hz. and insufficient photons (x-ray energy) result when even the highest capacity x-ray tubes are combined with realistically narrow x-ray beam slit widths and rapid scanning rates. In order to eliminate this problem a detector assembly has been proposed which enables the practice of energy subtraction radiography with the use of a constant output x-ray source. In accordance with this proposal, a dual layer planar dual energy radiation detector assembly has been suggested in the application EP-A-0 115 125 cited above. A first layer comprises a rectangular planar array of square detector photodiode elements including a first radiation sensitive scintillation material overlying the photodiodes and being selected for its primary response to radiation of a lower energy range. A second planar layer is located, or "stacked", directly behind the first layer, with respect to the x-ray tube, and comprises a similar rectangular array of detector elements congruent and aligned with the first layer. The second layer includes a second radiation sensitive scintillation material selected for its propensity to respond primarily to radiation of a higher energy level, which has passed through the first layer substantially without being detected. Such a dual energy detector structure, when used in conjunction with an x-ray tube emitting energy over a wide range, will provide data describing two separate images, i.e. one an image of lower energy x-radiatibn passing through the subject, the other being an image describing the pattern of higher energy radiation. A flat rectangular detector of any type, however, has the inherent disadvantage that radiation intensity falling upon the various detector elements is a function of the element's relative position in the array, in addition to being a function of the patient's body structure. This results from the fact that x-ray energy travelling from the x-ray tube focal spot toward a detector element near the image periphery must travel a longer path than x-ray energy directed toward the central portion of the detector. This disadvantage is a problem in both single and dual energy flat detector systems. With a flat detector, x-rays incident on centrally located detector elements enter the scintillator nearly normal to the receiving face. Rays incident on peripherally located elements enter the scintillators at angles which vary considerably from the normal. Under such conditions, a ray entering a peripheral element travels a longer path within its respective scintillator than does a ray entering central elements. Due to the longer paths travelled, a ray entering a peripheral element produces a greater scintillation than does a ray of equal value entering a central element, causing a response by the peripheral element that falsely exaggerates the indicated energy of the ray entering the peripheral element. This problem acquires yet another aspect where a dual energy detector is used having two sets of "stacked" detector element arrays. A greater proportion of a ray's energy is dissipated in passage through a peripheral element than in passage through a central element. Therefore, a smaller fraction of energy incident on a peripheral front layer detector element is left to actuate a rear peripheral detector element than would be left to actuate a rear central detector after passage through a front central element. This phenomenon distorts the relative responses of the dual energy detector arrays, or layers, in a way which is very difficult to correct. Where a flat dual energy detector, as described above, is used, there is an additional disadvantage. Due to parallax, image features viewed by the second detector layer, behind the first detector layer, will be seen as slightly spatially displaced with respect to the same image features as seen by the first detector layer. Therefore, the image produced by the second detector layer will differ slightly from that of the first, which difference will degrade the quality of the energy subtraction image which is derived when the first image data is subtracted from or combined with the second. It is an object of this invention to provide a radiographic system capable of producing shadow images but not computerised tomography images and having a dual energy detector wherein the imaging response is uniform over the detector, and which does not suffer disadvantages of parallax. According to the present invention there is provided a radiographic system capable of producing shadow images but not computerised tomography images comprising means arranged to pass a divergent beam of radiation from a source through a volumetric portion of a subject onto a radiation detector means to produce said shadow image of said portion of the subject, said detector means producing electrical output signals representative of the radiation incident thereon and the system further including imaging circuitry coupled to said detector means for utilising said output signals to produce said shadow image, and said detector means comprising a first detector structure comprising a material sensitive to radiation within a first energy range and having a first surface on which said radiation falls, and a second detector structure comprising a material sensitive to radiation within a second energy range and having a second surface on which said radiation falls and which overlies said first surface, characterised in that said first and second surfaces are curved about said source so as to eliminate substantially for each one of said surfaces differences in the lengths of the paths between said source and different points on that one of said surfaces. In one particular embodiment of the invention said beam is a divergent thin planar beam and said surfaces each define a portion of the surface of a cylinder whose axis is perpendicular to the plane of said beam. In such a system said source normally will have a focal spot from which said beam emanates and each said axis passes substantially through said focal spot. A system according to the invention suitably includes means for rotating said source relative to the subject so as to scan radiation across the subject, and means for simultaneously scanning said detector means along an arcuate path so as to detect the radiation from the source passing through the subject as the radiation scans across the subject. In another particular embodiment of the invention each said surface defines a portion of a surface of a sphere. In such a system said source normally will have a focal spot from which said beam emanates and each sphere has its centre substantially coincident with the focal spot. In US-A-4284895, DE-A-2831038 and DE-A-3311931 there are described radiation detector arrays which present a curved surface to radiation incident thereon from a radiation source so that the lengths of the paths between the source and different points on the surface are substantially the same. However, these arrays are used in computerised tomography (CT) radiographic systems, not in radiographic systems of the kind to which the present invention relates, i.e. which are specifically not capable of producing computerised tomography images. In DE-A-2831038 use in a radiographic system of multi-layered detectors to determine the energy profile of incident radiation for the purpose of effecting correction for radiation hardening in a subject is described, but these multi-layered detectors do not comprise a multi-energy imaging detector structure. Several radiographic system in accordance will now be described, by way of example, with reference to the accompanying drawings in which:- Figure 1 is an isometric view of a medical diagnostic x-ray system: Figure 2 is a block diagram illustrating components of the system of Figure 1; Figure 3 is an isometric view illustrating a first system in accordance with the invention; Figure 4 is an isometric view illustrating a second system in accordance with the invention: Figures 5 and 6 are isometric views showing detail of portions of two systems in accordance with the invention; Figure 7 is an isometric view illustrating detail of a portion of the system as shown in Figure 4; and Figure 8 is a graphical representation of a principle of operation of system embodiments shown in Figures 5 - 7. A system S for performing digital scan projection radiography (SPR) is illustrated in general form in Figure 1. The system S directs a pattern of x-rays through a patient P and produces, from information borne by the x-ray pattern emergent from, the patient's body, a representation, generally in the form of a visible image, describing internal structure or condition of the patient's body. The system S incorporates an x-ray source 10 for directing a beam of x-ray energy illustrated as a collection of rays 12 through the patient P and onto a detector assembly 14. A first collimator structure 16 defines a generally vertical fore slit 18 for collimating the x-rays emanating from the source into a spread beam lying generally within a vertical plane. A second collimator structure 20 defines an aft slit 22 located between the patient and the detector assembly, aligned with the fore slit and with the detector, for enhancing this collimation. Mechanical structure (not shown) maintains a mutually constant relative alignment between the collimators 16, 20, the x-ray source 10 and the detector assembly 14. In the preferred embodiment, mechanical means is provided for scanning the collimators and detector in unison relative to the patient's body in a manner described in more detail below. The mechanical structure can suitably comprise a gantry structure of known configuration (not shown) which physically holds the collimators and detector in a rigid alignment, and mechanical drive means to move the entire gantry to effect scanning. Alternately, the components can be coupled to individual drive mechanisms, and servo techniques can be employed in known fashion to maintain the desired alignment during scanning motion. In the preferred embodiment, mechanical scanner apparatus 24 is coupled to the detector assembly 14 to move the detector along a generally arcuate path defined by the arrows 26, 28. The arcuate path is centered about a vertical axis 34 through a focal spot 32 of the tube 10, described in more detail below. Pivoting apparatus 30 is coupled to the x-ray source. The apparatus 30 pivots the source, synchronously with detector and collimator arcuate motion, to continuously track the detector 14 and the mutually aligned collimators 16, 20. The x-ray source 10 comprises an x-ray tube, and associated power circuitry (not shown) for electrically actuating power circuitry (not shown) for electrically actuating the tube to produce x-rays (in pulsed or continuous mode) emanating from a focal spot 32 defined by the structure of the tube. Tube 10 produces x-rays by directing a stream of electrons onto an anode 33a of the tube, the anode rotating about an axis 33b. The pivoting motion effected by the pivot apparatus 30 causes the tube to pivot about the vertical axis 34 extending through the focal point 32. The axis 33b of the rotatable anode 33a is, in the Figure 1 embodiment, approximately parallel to the vertical axis 34 through the focal spot, the axis 34 being the axis of detector scanning as well, about which the arcuate detector path is centered. In some instances, it is desirable to tilt the axis 33b slightly, with respect to the axis 34, by about 8 degrees, or a similar small amount. This tilting optimizes the configuration of the pattern 12 of x-rays which emanate from the tube and propagate through the collimator slits 18, 22. More specifically, the tilt angle is chosen to maximise the uniformity of x-ray energy passing through all portions of the slits 18, 22. The tilting is desirable for uniformity maximisation because the x-ray energy produced by the tube 10 varies inherently with the angle 33c from the axis 33b at which energy emanates, and the degree of non-uniformity varies over the range of such angles. This is often called the "heel and toe effect" of the tube. The tilting angle is selected to enable the projection of energy through the slits 18, 22 from that range of angles 33c over which the x-rays are most uniform. It is believed preferable to couple the detector assembly 14 to the master drive of the scanner apparatus and to control the tube and collimators to follow, since detector positioning is more critical than tube positioning. An encoder 36 is coupled to the scanner apparatus 24 and produces a signal indicating the instantaneous position of the detector 14 along its arcuate path described by the arrows 26, 28. The output of the encoder 36 is directed to the pivot apparatus 30 for synchronising the pivoting motion of the x-ray tube 10 with the arcuate motion of the detector 14 and collimators 16, 20, to maintain continuous alignment between the x-ray beam, collimators and detector assembly during scanning motion. The scanner apparatus can be appropriately gated by a physiological signal, such as by an ECG signal 37, or by a signal indicating timing of administration of a contrast agent. Temporal subtraction studies can also be done, with sufficiently rapid retrace between scans. An example of a type of encoder apparatus is described in US Patent No. 4,015,129, issued on March 29, 1977 to Manring et al, and owned by the assignee of the present application. The encoder 36 may also be coupled to a current control 33 of the x-ray tube 10. The encoder can adjust the tube current and hence, the intensity of x-ray output, as a function of the location of the detector along its scanning path. In the embodiment described here, the tube current can be controlled to decrease as a function of the degree of detector displacement from the centre position along its scanning path. Thus, where the patient's body is less thick, i.e. near its right and left sides, x-ray output is reduced to maintain a more uniform x-ray flux at the detector throughout its scan. The detector assembly 14, a conventional detector array being shown in Figure 1, includes an array of individual detector elements, generally arranged within an elongated slot 28a defined by the detector assembly 14. The structure and arrangement of the detector elements is described in detail below. Each of the detector elements responds to light energy (generated by x-rays as described below) to produce an analog electrical charge signal which represents a characteristic of the x-ray which caused the production of the electrical signal. In operation, the detector, collimators and x-ray tube are moved to the left as in the direction ilustrated by the arrow 26 to prepare for a scan. In performing a scan, the x-ray tube 10 is actuated to produce x-ray energy. The scanner apparatus 24 and pivot apparatus 30 cooperate to synchronously scan the vertical spread beam of x-rays from left to right as shown in Figure 1 across the patient's body. During this scanning motion, the detector elements of the detector assembly 14 produce the analogue electrical signals. Analogue detector outputs from each of the detector elements are periodically sampled. Each sampling produces analogue signals representing a portion of image information. Over the course of the scan from one side to the other side, signals are developed describing a plurality of image lines, which together constitute an area image of the patient's internal body structure. The electrical signals are then digitised and processed to produce the desired patient imaging. Figure 2 illustrates a generalised block diagram of the system of Figure 1. In Figure 2, the x-ray source 10 directs x-rays to the detector assembly 14. The individual detector elements of the detector assembly 14 are mutually coupled to and associated with time delay and integrate (TDI) circuitry 43 which executes shift and add operations on the detector outputs in synchronism with detector scanning motion to produce image indicating analogue signals with enhanced signal-to-noise ratio. Signals from the detector 14 and TDI circuitry 43 are then transmitted to a data processor 44 which digitizes and processes the electrical signals. In response to commands from an operator's console 46, the data processor 44 produces various types of representations of internal body structure of the examined patient. In one mode, the data processor actuates a diagnostic viewing console 48 to produce directly a visible image of the patient's internal body structure which can be immediately employed by a radiologist for medical diagnostic purposes. In another mode, the data processor 44 stores digital information representing patient image data in one or more peripheral memories 50. Optionally, a camera 52 can be coupled to the data processor. Figure 3 shows one embodiment of a focussed detector in accordance with the invention. Figure 3 illustrates an x-ray tube 60 directing a spread beam of x-rays 62 through a patient P toward the detector arrangement 64. Fore and aft collimators 66, 68 defining aligned collimator slits are also illustrated. The remainder of the system is similar to that described above and is omitted for simplicity and clarity. The detector includes a first (front) array 70 of detector elements and a second (back) array 72 of detector elements located behind the first array with respect to the x-ray tube. Both the first and second arrays are aligned with the fore and aft slits. Radiation from the x-ray tube falls upon, and is partially absorbed by, the first array, and the remainder of the radiation, passing through the first array, falls upon and is detected by the second array. In this way, separate dual energy response is obtained, as explained in the Barnes published application cited above. Each of the arrays includes a single line of detector elements arranged along an arcuate path defined by a portion of a circle having its centre located at a focal spot 74 of the x-ray tube. The detector arrays scan along arcuate paths concentric with the path indicated by the arrows 69. The concentric paths are centered about a vertical scanning axis 75 extending through the tube focal spot 74. The axis 77 is approximately perpendicular to the scanning axis 75. This geometry reduces the non-uniformity of the x-ray energy across the beam set 62 propagating through the collimator 66, 68 by eliminating the effect of the "heel and toe effect" characteristic of the x-ray tube. According to the heel and toe effect, the x-ray energy from a tube varies as a function of the angle of x-ray emission with respect to the axis of anode rotation. Such radiation is far less a function of the angle of the x-ray propagation taken radially with respect to the anode rotative axis. In the Figure 3 embodiment, the axis 77 is also tilted slightly from the perpendicular to the approximate plane defined by the aligned collimators 66, 68 and detector array layers 70, 72. The amount of tilting is about 8 degrees from the perpendicular to the plane so defined. The tilting is desirable for reasons analogous to thos explained with respect to the tilting in Figure 1. Each detector element comprises a photodiode. Overlying each photodiode is a scintillation material responsive to x-rays to produce visible light energy. The scintillation material used in connection with the first array differs from that used in connection with the second array. The scintillation material associated with the first array is selected for its ability to absorb and produce light in response to x-rays from the source falling primarily within a relatively low energy range. The higher energy x-rays pass through the first array and fall on the second array, causing the scintillation material associated with the second array to produce light which is detected by the individual detector photodiodes of the second array. Suitable types, thicknesses and physical configurations of the scintillation material are defined in the above cited Barnes published European patent application. The preferred detector embodiments of this invention described can suitably include a radiation filter, made of copper or brass sheet or other similar material, located at the region indicated by reference character 73, to "harden" the radiation energy reaching the second detector layer or array. Such filter elements are used to improve energy discrimination between the two detector layers in ways as are described in the above cited Barnes published European patent application. The scintillation or phosphor material used in conjunction with the individual photodiode arrays discussed herein can suitably comprise a uniform single portion or layer of phosphor material overlying the entirety of the photodiode array without breaks or interruptions in its surface. Thus, the configuration of this embodiment forms a uniform smooth curved surface which comprises the energy receiving face of the detector array. In another embodiment, each detector element photodiode can be provided with its own individual portion of scintillation phosphor material, rather than the entire array being covered with a single piece of such material. In this latter embodiment, the individual receiving faces of each of the detector elements, with their scintillators, collectively form a curved surface which is approximately smooth. Known electronic and/or software correction means can be used to compensate, if need be, for any non-uniformity of transmission characteristics through the front detector array. This correction means can be associated with data processor 44. Figure 4 illustrates another embodiment of the invention in the form of an alternate detector array configuration. In Figure 4. energy from an x-ray tube 80 is directed through a collimator (not shown) and in an area beam 82 through the patient. emerging therefrom to all in a relatively large area of a detector assembly 84. In the embodiment of Figure 4, a first detector array 86 includes a relatively large number of individual detector elements arranged in an area pattern, with the receiving surface of the first detector array collectively defining a portion of a sphere having its centre at a focal spot 87 of the x-ray tube. The detector element arrangement (not shown) is as described in the above cited Kurger patent. A second similar detector array 88 is located behind the first array with respect to the x-ray tube. The second array has a receiving surface defined by a portion of a sphere having a radius slightly larger than the sphere referred to in connection with the first array, and whose centre is also located at the x-ray tube focal spot. The source/detector arrangement shown in Figure 4 incorporating the area beam can, but need not, be scanned. That is, the area beam can be directed constantly or in pulsed mode to pass simultaneously through a predetermined relatively large area of the patient's body, or the beam and detector can be scanned in unison,analogous to the manner of scanning described in connection with Figure 1. An alternative embodiment employs the spherically or cylindrically configured, detector arrays of Figure 4 and in Figures 5-7 (described below) in conjunction with the thin spread beam and collimator arrangement as illustrated in Figure 3. In this alternate embodiment, the spread beam is scanned across the area detector, the detector remaining stationary relative to the patient. Techniques for processing image data from non-scanned detector arrays, or single line scanned arrays in digital radiography are well known. Data processing of scanned multi-linear arrays is discussed below. Other embodiments of focussed detector configurations in accordance with the invention are illustrated in Figures 5 and 6. Figure 5 shows a portion of each of two stacked detector arrays 92, 94. The detector of Figure 5 comprises a multiplicity of detector elements 96 whose receiving surfaces collectively define approximately a portion of a cylinder having a horizontal axis intersecting the focal spot of the x-ray tube. Either a single layer, or a dual layer, or "stacked" detector, as shown, can be used. Figure 6 illustrates another embodiment. wherein the detector element receiving faces 95 define collectively a portion of a cylinder having a vertical axis through the x-ray tube focal spot. A dual layer (97,99) arrangement is shown, but a single layer can be used if desired. It is important to note that in the detector arrangement of Figures 4-7, (Figure 7 to be discussed below) known time delay and integration (TDI) circuitry can be employed, where the beam and detector are synchronously scanned, to enhance the signal-to-noise ratio of information derived from the scanned detector. In the embodiment of Figure 4, for example, a rectangular array of detector elements can be used. When such an array is used, the techniques for employing TDI output signal enhancement are described in the above cited Kruger patent. The embodiments of Figures 5-7 can also incorporate rectangular arrays of elements and TDI. A further aspect of this invention is illustrated in Figures 5-7. Figure 7, for example, depicts a portion of a double layer (101, 103) spherical (three dimensionally curved) detector assembly 104, similar to that of Figure 4, but wherein individual detector elements are arranged in a staggered fashion. Figures 5 and 6 illustrate embodiments wherein the detector elements are arranged in staggered patterns and collectively define a portion of a cylinder. In the embodiments of Figures 5-7, incorporating staggered arrays of detector elements, time delay and integrate (TDI) signal enhancement techniques can also be employed. Each column (extending substantially vertically in Figures 5-7) is offset vertically from its adjacent columns by a distance equal to one half the centre-to centre spacing between adjacent elements along a column. It should be understood that, while the preferred embodiments described herein are intended to include a dual layered stacked detector, the invention also encompasses embodiments similar to those of the Figures, but employing more than two stacked detector layers used analogously to those described here. Also, it should be understood that each of the embodiments of Figures 4-7 preferably include a beam hardening filter analogous to the filter 73 shown in Figure 3, and the x-ray tube geometry described in connection with Figures 1 and 3. The relationships among the signals appearing in the various TDI sampling periods and among the various detector elements, in implementing the TDI technique in conjunction with the staggered array. are represented by the following relation: The total accumulated charge corresponding to the j-th pixel of the i-th row, i.e., P (i j), of the total derived image is given by the following expression: where: d(m,n,k) = the charge accumulated in the n-th detector element of the m-th row during the k-th sampling period, and N = the number of detector elements per row. See Figure 8 for a graphical representation of the notation of the above equation. It is to be understood that the foregoing description of the present invention is intended as illustrative, rather than exhaustive, of the invention.

Claims: A radiographic system capable of producing shadow images but not computerised tomography images comprising means (S) arranged to pass a divergent beam of radiation (62 or 82) from a source (60 or 80) through a volumetric portion of a subject (P) onto a radiation detector means (64 or 84) to produce said shadow image of said portion of the subject, said detector means (64 or 84) producing electrical output signals representative of the radiation incident thereon and the system further including imaging circuitry (43, 44, 48) coupled to said detector means (64 or 84) for utilising said output signals to produce said shadow image, and said detector means (64 or 84) comprising a first detector structure (72 or 88) comprising a material sensitive to radiation within a first energy range and having a first surface on which said radiation falls, and a second detector structure (70 or 86) comprising a material sensitive to radiation within a second energy range and having a second surface on which said radiation falls and which overlies said first surface, (P) characterised in that said first and second surfaces (70, 72 or 86, 88) are curved about said source (60 or 80) so as to eliminate substantially for each one of said surfaces (70, 72 or 86, 88) differences in the lengths of the paths between said source (60 or 80) and different points on that one of said surfaces (70, 72 or 86, 88). A system according to Claim 1 in which said source (60 or 80) is an x-ray tube (60 or 80) and said surface materials sensitive to radiation comprise x-ray sensitive material. A system according to Claim 1 or Claim 2 wherein said beam is a divergent thin planar beam (62) and said surfaces (70 or 72) each define a portion of the surface of a cylinder whose axis is perpendicular to the plane of said beam (62). A system according to Claim 3 wherein said source (60) has a focal spot (74) from which said beam (62) emanates and each said axis passes substantially through said focal spot (74). A system according to any one of the preceding claims including means (30) for rotating said source (60) relative to the subject (P) so as to scan radiation across the subject (P), and means (24) for simultaneously scanning said detector means (64) along an arcuate path (69) so as to detect the radiation from the source (60) passing through the subject (P) as the radiation scans across the subject (P). A system according to Claim 1 or Claim 2 in which each said surface (86 or 88) defines a portion of a surface of a sphere. A system according to Claim 6 in which said source (80) has a focal spot (87) from which said beam (82) emanates and each sphere has its centre substantially coincident with the focal spot (87). A system according to any one of the preceding claims in which each of said detector structures (70, 72 or 86, 88) comprises a plurality of individual detector elements (96) each having a radiation sensitive face (95), the faces (95) collectively defining said first or second surface. A system according to Claim 8 in which the source (60) has a focal spot (74) and each detector element (96) of one of the detector structures (72 or 84, 70 or 86) is aligned with a corresponding detector element (96) in the other of the detector structures (70, 72 or 86, 88) along a line extending from the focal spot (74). A system according to any one of the preceding claims in which the imaging circuitry (43, 44, 48) includes time delay and integrate circuitry (43).

Inventors: SONES RICHARD A, SONES, RICHARD A.
Applicants: PICKER INT INC, PICKER INTERNATIONAL, INC.
Application year: 1985
Main classification: A61B6
IPC codes: G01T1, G01N23, A61B6