EP0000079B1 - Röntgenabtastsystem und Verfahren. - Google Patents

Röntgenabtastsystem und Verfahren. Download PDF

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Publication number
EP0000079B1
EP0000079B1 EP78200018A EP78200018A EP0000079B1 EP 0000079 B1 EP0000079 B1 EP 0000079B1 EP 78200018 A EP78200018 A EP 78200018A EP 78200018 A EP78200018 A EP 78200018A EP 0000079 B1 EP0000079 B1 EP 0000079B1
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EP
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Prior art keywords
ray
detector
scanning
anode plate
target anode
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EP78200018A
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English (en)
French (fr)
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EP0000079A1 (de
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Richard David Albert
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Individual
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    • HELECTRICITY
    • H05ELECTRIC TECHNIQUES NOT OTHERWISE PROVIDED FOR
    • H05GX-RAY TECHNIQUE
    • H05G1/00X-ray apparatus involving X-ray tubes; Circuits therefor
    • H05G1/08Electrical details
    • H05G1/66Circuit arrangements for X-ray tubes with target movable relatively to the anode
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61BDIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/40Arrangements for generating radiation specially adapted for radiation diagnosis
    • A61B6/4021Arrangements for generating radiation specially adapted for radiation diagnosis involving movement of the focal spot
    • A61B6/4028Arrangements for generating radiation specially adapted for radiation diagnosis involving movement of the focal spot resulting in acquisition of views from substantially different positions, e.g. EBCT
    • AHUMAN NECESSITIES
    • A61MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61BDIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/50Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment specially adapted for specific body parts; specially adapted for specific clinical applications
    • A61B6/51Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment specially adapted for specific body parts; specially adapted for specific clinical applications for dentistry
    • HELECTRICITY
    • H01ELECTRIC ELEMENTS
    • H01JELECTRIC DISCHARGE TUBES OR DISCHARGE LAMPS
    • H01J35/00X-ray tubes
    • H01J35/24Tubes wherein the point of impact of the cathode ray on the anode or anticathode is movable relative to the surface thereof
    • H01J35/30Tubes wherein the point of impact of the cathode ray on the anode or anticathode is movable relative to the surface thereof by deflection of the cathode ray

Definitions

  • This invention relates to radiographic apparatus and procedures and more particularly to scanning X-ray systems which produce signals which may be used to present a visible image at a cathode ray tube screen or the like.
  • the present invention was initially developed for usage in dental and medical radiology, and to facilitate description the invention will be herein discussed with reference to this particular field of use. As will be apparent, the apparatus may also be advantageously adapted to various other radiographic operations.
  • an X-ray film packet is inserted into the patient's mouth near the teeth or other anatomical structure to be imaged.
  • the procedure makes use of an X-ray tube of the form which generates X-rays at a fixed point on an anode and in which the X-rays radiate out from that fixed point.
  • the tube is provided with a shield cone which is directed at the film packet through the teeth or other structures to be imaged. Precise positioning of both the film packet and the shield cone is necessary to obtain useful X-ray images.
  • Another very serious problem is that undesirably large radiation dosage exposures of the patient are needed in order to produce a complete set of dental X-ray images. This is in part a result of the very low detection efficiency of the unscreened X-ray film commonly used for dental X-ray operations.
  • panoramic or wide angle X-ray images are produced by generating a narrow linear X-ray beam which is revolved during the exposure about an axis of rotation situated within the patient's head.
  • the X-ray tube is essentially a conventional one at which X-rays are generated at a small fixed point on an anode.
  • Radiation generated at this point is collimated by a first slit which is parallel to the axis of rotation and then passes through the patient's head and then through a second similar collimating slit situated in front of a screened film cassette which is rotated in synchronism with the rotational movement of the X-ray beam.
  • the tube and detector motion causes the X-ray beam to sweep across the intervening anatomical structures.
  • a panoramic two-dimensional strip image is produced of curved anatomical structures in the patient's head such as the mandible or maxilla.
  • the conventional pantomographic image technique is itself subject to several disadvantages. It is necessary that the X-ray beam pass through the entire skull of the patient, even if it is only desired to obtain an image of a portion of the skull such as the dental arch. Consequently, unwanted images are superimposed upon the desired image data. This makes interpretation of the image more difficult and detracts from the general quality of the image by obscuring desired data to some extent with undesired information. Moreover, radiation exposure remains undesirably high as the X-ray beam must necessarily penetrate through the entire skull.
  • Anatomical structures which are not of particular interest are thereby necessarily subjected to radiation dosage which does not contribute any useful information but instead detracts from the quality of the desired data. Further a significant amount of X-ray scattering occurs during passage of the X-ray beam through the patient's entire head creating a background fog in the image on the developed film which undesirably limits the range of contrast in the image and which may cause loss of definition.
  • the general form of scanning X-ray system disclosed in the above-identified prior patent and copending applications dispenses with the use of film as an X-ray detection medium and produces signals which may be used to produce a visible image on the screen of a cathode ray tube display device including instantaneous images if desired. Radiation dosage of the patient is substantially reduced.
  • the system may be utilized to image only a selected portion of a subject such as a patient's dental arch for example without including superimposed data from other regions of the subject.
  • Image data may be electronically stored on magnetic tape or by any of various other data storage means and the image data may also readily be processed by various electronic enhancement techniques to further improve image quality or to emphasize specific image characteristics.
  • a system of the general type described in the above-described patent and copending applications uses a scanning X-ray tube in which an electron beam is systematically swept in a raster pattern on a broad target or anode plate to produce a moving point source of X-rays.
  • the region of the subject which is to be imaged is situated between the anode plate of the X-ray tube and an X-ray detector which is small in relation to the size of the raster pattern and which may therefore readily be situated in the oral cavity or the like of a dental or medical patient or in similarly constricted interior spaces of an inanimate subject.
  • the raster sweep signals of a cathode ray tube display are coordinated with the scanning action of the electron beam in the X-ray tube and a signal derived from the X-ray detector output is applied to the intensity signal terminal of the cathode ray tube.
  • a visible radiographic image of the region of the subject situated between the X-ray source and the detector is produced on the screen of the display device.
  • a scanning X-ray system of this general type should possess certain specific capabilities.
  • radiation dosage should be minimized to the extent possible while producing an image of high definition and contrast range.
  • the dentist or other operator should be able to position the X-ray detector very precisely relative to the X-ray tube at any of a plurality of different positions within the patient's mouth, or in other constricted spaces, with a minimum of difficulty and with maximum patient comfort.
  • the effective focal length of the system be readily and precisely changeable in order to obtain images of different degrees of magnification.
  • Such a system should minimize optical distortions and other forms of image degradation, which can be present in apparatus of this general form, in order to facilitate image interpretation.
  • a scanning X-ray apparatus for producing radiographic image date comprising a housing in which an X-ray tube including a vacuum enclosure forms a vacuum region to enclose a target anode plate and an electron gun for producing an electron beam, said X-ray tube further having electron beam deflector means for directing said electron beam to successively different points on the surface of said target anode plate so as to produce a moving X-ray origin point at said target anode plate, and associated with said X-ray tube is a probe member supporting an X-ray detector possessing an active area smaller than the range of motion of said moving X-ray origin point so as to detect X-rays traversing an object disposed between said target anode plate and X-ray detector, the base portion of said probe member being remote from said X-ray detector, said probe member including means for transmitting X-ray count data from said detector to said base portion of said probe member, receiving means including at least one output for transmitting electrical signals indicative of said X-ray counts,
  • a scanning X-ray system 11 which greatly facilitates the obtaining of a variety of different forms of dental radiograph including the providing of instantaneous high-quality images with relatively low radiation dosage of the patient.
  • the apparatus is depicted in Figure 1 as utilized to produce a pantomographic image of the left half of the lower dental arch of a subject 12.
  • the same apparatus may then be quickly and conveniently adjusted to provide additional images for completing a set of pantomographic images, to provide one or more periapical images of individual teeth of particular interest and may also easily be used to produce images of other anatomical structures or of inanimate objects.
  • Salient components of the scanning X-ray system 11 include a scanning X-ray tube 13 of specialized construction, an X-ray detector probe 14A extending outwardly from the face of the tube and being supported thereby and an electrical control and signal processing circuit 16 including a cathode ray tube or other type of X-Y display device 17 of the form having a screen 18 at which visible images are displayed in response to X and Y axis sweep frequency signals received at terminals X and Y respectively and in response to intensity signals received at another terminal Z.
  • Oscilloscopes, television receivers and the like of known construction may readily be utilized as the display device 17.
  • X-ray tube 13 includes a vacuum enclosure 19, formed of glass or other suitable electrically insulative material, which defines an evacuated region 21.
  • Vacuum enclosure 19 has a relatively narrow end containing an electron gun 22 of suitable known construction for producing and accelerating an electron beam 23 towards an opposite target end of the enclosure.
  • the target end of enclosure 19 is larger than the electron gun end and is preferably of rectangular configuration where the system is to be used primarily for dental X-ray operations.
  • the target end of vacuum enclosure 19 is formed at least in part by a target anode plate 24 having at least an inner surface 26 consisting of one of the electrically conductive metals which produce X-rays upon being bombarded by high-energy electrons, copper, tungsten and tantalum being examples of metals suitable for the target anode surface 26. Contrast in the image is enhanced if the X-rays which reach the detector are monochromatic or nearly so. This may be arranged for by an appropriate selection of the thickness of the target anode plate 24 since elemental target materials tend to be more transmissive of their own characteristic X-rays than of other X-ray wavelengths.
  • Vacuum enclosure 19 is disposed within a housing 27 to support and protect the enclosure and to enable the mounting of additional components on the tube as will hereinafter be described.
  • Housing 27 may be formed of radiation-absorbent material such as steel or of any of various known plastics which contain sizeable admixture of a heavy radiation-absorbent metal or, as in this example, of plastic having a layer 27' of radiation-absorber at the outer surface.
  • housing 27 has a rectangular collimator receiving opening 28 in the region of the face of the tube which opening conforms in size and configuration with the target anode plate 26 of the tube.
  • An X-ray collimator 29A which will hereinafter be discussed in more detail, is received and supported in opening 28.
  • electron beam 23 isswep in a raster pattern on target anode plate 26.
  • the electron beam movement includes a repetitive sweep-movement back and forth across the anode surface 26 in a first direction which is parallel to the plane of Figure 1, and which is herein ⁇ designated the X-axis direction.
  • the electron beam is also swept at a slower rate back and forth in an orthogonal direction which in this example is normal to the plane of Figure 1 and is herein designated the Y-axis direction.
  • the combining of these two movements causes the electron beam to sweep successively along a series of substantially parallel lines which jointly define a rectangular raster pattern area on surface 26 of the anode plate.
  • Magnet deflector 31 may be of the form having four magnetic poles angularly spaced in quadrature, of which a single pole 32Y and winding 33Y appears in Figure 1, and having an annular ferromagnetic yoke 34 which encircles each of the poles.
  • Deflector 31 is disposed coaxially around the vacuum enclosure 19 between the electron gun 22 and target anode plate 26 within an annular groove 36 in the inside surface of housing 27.
  • the focal length of the scanning X-ray system is selectively changed by making certain adjustments which will hereinafter be explained in more detail.
  • One adjustment which compensates for an optical distortion which could otherwise occur upon a change of focal length requires the providing of means for selectively shifting the axial position of the area in which beam deflection occurs so that the beam deflection area may be moved further from target anode plate 26 or closer to the target anode place.
  • the annular groove 36 in which deflector 31 is situated is of greater length along the axis of the tube than is the deflector itself so that the deflector may be moved in the axial direction towards the target anode plate 26 or toward the electron gun 32 as desired.
  • tabs 37 extend radially outwardly from opposite sides of the deflector through slots 38 provided in housing 27 for that purpose.
  • a boss 39 is formed on housing 27 adjacent the forward end of each slot 38 and one of a pair of disengageable screws 41 extends through each tab 37, through a series of annular washers 42 and engages in a threaded bore 43 in the adjacent boss 39.
  • the deflec-, tor 31 may be located and fixed at any desired axial position within, the limits established by the length of groove 36.
  • more quickly operated axial positioning means of any of various forms may be utilized in place of the screws 41, the threaded rotatable telescoping sleeve mechanisms commonly used for changing the focal length of photographic cameras by axial movement of lenses being one example.
  • this example of the invention utilizes beam of deflection means of the magnetic variety, electrostatic beam deflection may also be used.
  • X-rays produced at a target anode plate surface 26 and transmitted through collimator 29A converge at a relatively small X-ray detector 44 situated at the distal end of the probe 14A at a location on the central longitudinal axis of the X-ray tube.
  • the location of detector 44 is spaced from the face of the tube, including collimator 29A, to enable positioning of the subject 12 which is to be imaged between the collimator and detector.
  • X-rays received at the detector 44 must first pass through the anatomical region or the like which is to be imaged.
  • detector 44 should preferably have a radiation-sensitive area as small as possible consistent with the need to obtain an adequate count rate from the amount of radiation which is received.
  • scintillation crystal detector 44 is necessarily depicted in Figure 1 as being somewhat larger than is usually optimum in actual practice, crystals measuring less than one millimeter in size being typical. Scintillation crystals of this kind respond to individual X-rays by producing scintillations of visible light which may be converted to electrical X-ray count signals by photosensitive means.
  • detector scintillator 44 preferably is formed in the shape of a sector of a sphere and is situated coaxially with the axis of the X-ray tube with the apex of the sector being most distance from the X-ray tube and being at the point of convergence of the X-rays from target anode plate 26 that are transmitted through collimator 29A.
  • the probe 14A In addition to supporting and positioning the X-ray detector 44, the probe 14A also transmits the X-ray count signals produced by the detector to receiving means which in this example includes a photomultiplier tube 46 of known internal construction which in this example is situated adjacent the small end of vacuum enclosure 19 within housing 27, the housing having a protruding portion 47 formed to receive the photomultiplier tube.
  • the probe 14A is primarily formed of a light pipe core 48 of light-transmissive material, the detector 44 scintillation crystal being partially embedded in or otherwise optically coupled to the light pipe core 48 near the end of the core which is remote from the X-ray tube.
  • a coating 49 of opaque material encloses all surfaces of the core 48 and of the X-ray detector 44 which would otherwise be exposed to ambient light.
  • an additional inner lining 49 of lead or other highly radiation-absorbent material may enclose the end portion of the core 48 in the region of detector 44 except for the surface of the detector which faces the X-ray tube.
  • Probe 14A has a base end 51 A releasably securable in any selected one of a series of attachment means 52A, 52B, 52C and 52D situated at different positions on the X-ray tube.
  • attachment means 52A in this example is situated at one side of the collimator-receiving opening 28 at the mid-plane of the X-ray tube while a second essentially similar attachment means 52B is provided at the same plane on the opposite side of the collimator-receiving opening.
  • Another attachment means 52C is located below the center of the collimator-receiving opening 28 while still another such attachment means 52D is situated above the center of the opening.
  • Attachment means 52A may include a cylindrical passage 53 penetrating for a distance into the face of housing 27 and shaped to receive base portion 51 A of probe 14A.
  • a linear rib or key 54 is formed on base portion 51 A of the probe and is received in a conforming axially extending slot 56 in the wall of passage 53.
  • detent means may be provided which in this example consist of a plunger 56 axially slidable in a passage 57 at right angles to passage 53 and having a spherical end surface suitable for engaging a conforming spherical cavity in the side wall of the base portion 51 A of the probe 14A.
  • Plunger 56 is biased towards passage 53 by a compression spring 58 and snaps into detenting position when the probe has been fully inserted while enabling easy removal of the probe when it is to be replaced with a probe of different configuration or when it is to be shifted to a different one of the attachment means 52.
  • the end 59 of the base portion 51 A of the probe 14A is a projecting end of the light pipe core 48 having a conical configuration.
  • Each of the other attachment means 52B, 52C and 52D may have a similar construction except insofar as each such attachment means has a different angular orientation with the keyway slots 56 of each attachment means always being adjacent the collimator-receiving opening 28 and the detent means always being on the outward side.
  • a light pipe 61 is disposed within housing 27 and has a broad base portion 62 disposed against the photosensitive surface of the photo tube 46. Forward from base portion 62, the light pipe 61 divides into four arms 63A, 63B, 63C and 63D. Arm 63A has a forward end extending into the inner end of passage 53 of attachment means 52A and has a conical indentation receiving the pointed end 59 of the base 51 A of probe 14A.
  • the juncture between end 59 of the probe and the forward end of arm 63A of light pipe 61 may be coated with silicone grease or other known materials suitable for such purpose.
  • the other arms 63B, 63C and 63D of the light pipe 61 extend to the other attachment means 52B, 52C and 52D respectively in an essentially similar manner.
  • opaque closure plugs 64 are inserted in the passages 53 of such attachment means and are disengageably retained therein by the same detent plungers 56B which otherwise engage a probe.
  • the probe 14A depicted in Figures 1 and 2 is designed to facilitate the obtaining of pantomographic images embracing a sizable portion of a dental arch of a subject 12. Such images are usually a panoramic image of approximately one-half of the dental arch. For such purposes, it is usually preferable to locate the X-ray detector 44 in the general vicinity of the patient's third molar tooth at the opposite side of the dental arch. To effect this detector disposition with a minimum of discomfort to the patient this particular probe 14A has an essentially quarter circular configuration between the linear base portion 51 A and the immediate region of the detector 44 so that it curves around the front of the patient's face and extends a small distance into the opposite side of the mouth.
  • closure 64 is removed from attachment means 56B and the probe 14A is removed from attachment means 52A.
  • the probe is then rotated and inserted into attachment means 52B, and the closure 64 is then inserted into attachment means 52A.
  • the probe 14A then extends from the opposite side of the face of the X-ray tube 13 as depicted by dashed lines 14A' in Figure 1, the X-ray detector 44 again being situated at the same centered position as before.
  • the X-ray tube 13, including the probe 14A may then be rotated around to the opposite side of the patient or alternately the patient may be turned to locate the detector 44 in an essentially similar manner but at the other rear molar 65L of the dental arch.
  • the ability of a scanning X-ray system 11 of this general form to produce a radiographic image of a subject is not dependent on the presence of a collimator 29A since only X-rays which are directed towards the minute detector 44 can contribute to the image.
  • the primary function of the collimator 29A is to minimize radiation dosage of the subject.
  • the collimator acts to enhance contrast in the image by reducing spurious counts at detector 44 from scattered radiation, X-ray fluorescence and the like.
  • the collimator 29A may detract somewhat from definition in the image as compared with a similar scanning X-ray system not having a collimator but the magnitude of this effect can easily be kept within acceptable limits and may be made insignificant since any reduction of definition caused by the collimator may be made to be less than the inherent definition limitations of the image display device 17.
  • the matter of reducing patient radiation dosage is generally a more critical one, it is usually desirable in dental or medical usages to operate system 11 with the collimator 29A in place.
  • the collimator 29A of this invention is of a specialized construction which provides for a high degree of reduction of patient radiation exposure while minimizing any consequent loss of image definition.
  • the collimator 29A includes a collimation element 67 formed of lead glass or other radiation absorbent material having similar properties.
  • Collimation element 67 is formed with a large number of spaced-apart radiation transmissive passages 68, of very small cross-sectional area, for transmitting X-rays which originate at target surface 26 towards the detector 44.
  • the passages 68 are aligned in directions which are convergent at the X-ray detector 44. In other words, each of the passages 68 extends along a separate radium of a hypothetical sphere having the apex end of detector 44 as a center.
  • the impact of the electron beam 23 on target anode place 26 causes X-rays 69 to be emitted in all directions from the point of impact.
  • the effect of the collimator 29A is to absorb those of the X-rays 69 which are emitted in the general direction of the patient but which are not traveling precisely towards the detector 44 and therefore cannot contribute useful information to the desired image.
  • the collimator has a similar effect on secondary X-rays which may be produced by interaction of a primary X-ray 69 with material behind or within the collimator.
  • the radiation transmissive passages 68 are convergent toward detector 44 when viewed in an orthogonal plane as well as in the plane of Figures 1 and 3 and thus the effect of the collimator is to intercept X-rays 69 which are directed towards the subject 12 other than the particular X-rays 69' which are directed exactly towards detector 44, regardless of the angular direction in which the unusable X-rays 69 deviate from a line extending towards the detector.
  • the X-rays produced at target anode plate 24 by impact of the electron beam 23 include X-rays of various different energies and the lowest energy component of the X-ray spectrum is of little or minimal value insofar as the production of the desired image is concerned.
  • one or more absorbent filtering means formed of a low-atomic-number material, such as aluminum for example may be situated between the target anode plate 26 and the patient.
  • a first such filter layer 71 which may typically be two or three millimetres of aluminum or equivalent radiation absorber, is disposed against the forward surface of the target anode plate.
  • Additional suppression of low-energy X-rays may be provided for by disposing another layer 72 of filter material on the surface of collimator element 67 which is closest to the target anode plate and by providing still another layer of such material 73 on the outer surface of the collimator element.
  • the outer layer 73 of filter material has the further beneficial effect of absorbing scattered X-rays arising from scattering effects within the collimator 29A itself.
  • the radiation transmissive passages 68 may be filled with a material such as aluminum or any of various suitable plastics which absorb very low-energy X-rays in order to further reduce radiation dosage from X-ray scattering within the collimator.
  • the radiation-transmissive passages 68 are not necessarily unobstructed open passages nor are they necessarily transmissive to all forms of X-rays.
  • the term radiation-transmissive passage is herein used to define a zone through the collimator which is transmissive to X-rays of a desired energy but not necessarily transmissive to other things..
  • a collimator element 67 having the physical characteristics described above may be manufactured by very precise drilling of the passages 68 through a lead glass element and if formed in this manner, laser beam drilling techniques may be preferred in order to enable the providing of a very large number of closely spaced passages 68 of a very minute cross-sectional area..As a practical matter it is preferable to form the collimator element 67 by fiber optical techniques similar to those heretofore used for manufacturing microchannel plates as used for example in night-vision devices and image- intensifiers. in one such technique, for example, the element 67 is initially formed by disposing a large number of tubular lead glass rods in parallel relationship with the passages through the rods being initially filled with a water- or acid-soluble core material.
  • the array of glass rods is then fused by heating and is then drawn, in which process interstices between the adjacent glass rods are eliminated and cross-sectional size is reduced.
  • the fused array known as a boule, is then cut transversely to produce an element of the desired thickness.
  • the resulting fused array is then etched to remove the soluble core material from the passages and the faces are polished to form the desired microchannel plate.
  • Elements can be produced by these techniques in which the passages 68 are typically about 25 microns in diameter and have a center-to-center spacing of about 37 microns. After processing to the extent described above, the collimating element has passages 68 which are parallel rather than convergent.
  • a sagging process may be used in which the microchannel plate is disposed on a sphere having a center at the point towards which the passages are to converge.
  • the element is then heated to a temperature where sagging occurs either as a result of the weight of the mass of the microchannel plate or if necessary with the aid of externally applied pressure.
  • This sagging causes the plate to conform to the surface of the sphere which deformation has the effect of causing the initially parallel passages 68 to now be convergent at the center of curvature of the sphere.
  • the primary effect of this sagging is to produce the desired convergence of the radiation-transmissive passages 68.
  • the spherical curvature of the collimator element 67 as a whole as depicted in Figure 1 is also produced by the sagging operation and may have an advantage in many cases in that it enables closer fitting of the X-ray source against curved exterior portions of the patient's anatomy such as the curvature in the jaw region.
  • the collimator element 67 need not have an overall curvature but can be planar provided that the internal passages 68 have the desired. convergence, an example of such a flat collimator being hereinafter described.
  • the collimator may be produced by a somewhat different process which does not involve the sagging step.
  • a fused array of hollow-lead glass capillary tubing may be heated and drawn into the form of a long cone both to reduce passage size and to produce the desired convergence of all passages towards a single point.
  • a desired flat section of the cone may then be produced by transverse cuts, with the resulting flat end surfaces being polished to produce. the finished collimator element.
  • This technique is capable of producing passages 68 down to around 200 microns in diameter without difficulty although with care still smaller passages may be produced.
  • the embodiment of Figure 1 utilizes a releasable collimator 29A so that the collimator may be removed and replaced with another collimator in which the passages 68 converge at a point further out from or closer to the target anode plate depending on whether focal length is to be increased or decreased.
  • the collimator element 67 is disposed in a rectangular frame 74 which fits into the collimator-receiving opening 28 at the face of the tube and which is disengageably retained therein by threaded screws 76 and which is preferably formed of radiation-absorbent material.
  • screws 76 extend into the forward side portions of housing 27 and enter a short distance into appropriately located bores in the side of the collimator frame 74.
  • the collimator 29A may be removed from opening 28 and another collimator of different focal length may be inserted and may be held in place by re-engaging the screws.
  • Screws 76 may be replaced by more elaborate but more quickly operated latching means in instances where frequent changes of focal length are contemplated.
  • a pantomographic image of the left side of the dental arch of a subject 12 will be produced on screen 18 of X-Y display device 17 by coupling the output of an X-sweep frequency generator or oscillator 77 to both the X-sweep terminal 78 of beam deflector 31 and to the X-sweep terminal 79 of the display 17.
  • the output of a Y-sweep frequency generator 81 which produces a repetitive sweep wave form of substantially lowerfre- quency than that of the X-sweep frequency generator 77, is coupled to the Y-sweep signal terminal 82 of the beam deflector 31 and also to the Y-sweep terminal 83 of the display device 17. Accordingly, the electron beam 23 is caused to repetitively sweep through a raster pattern on target anode plate 24 and the electron beam sweep of the display device 17 is coordinated with that of the source. Circuits for this purpose will hereinafter be described in more detail.
  • the X-ray count signals produced at detector 44 vary in number in the course of this scanning action in accordance with variations of the radiation-transmissiveness of radiolucence of the particular portion of the patient's anatomy which is being intercepted by X-rays 69' at each successive time in the course of the scanning action.
  • the count signals which are initially optical signals, are transmitted by light pipe core 48 of probe 14A and then by the internal light pipe 42 of the X-ray tube to photomultiplier tube 46 where the count signal data is translated into an electrical voltage signal.
  • the output signal terminal 84 of photomultiplier tube 46 is coupled to the Z or intensity signal terminal 86 of display device 17 through a video type of amplifier 87.
  • successive points in the visible image at the screen 18 of device 17 have a degree of illumination-intensity determined by the radiation transmissiveness at the corresponding points in the region of the patient's anatomy which is being scanned and the image at screen 18 constitutes the desired radiographic image.
  • the image may be viewed directly at screen 18.
  • the image may also be viewed in that manner, without regard to such scanning time limitation, if a long persistence cathode ray tube screen 18 is used and preferably one of the adjustable persistence type.
  • a camera 18' may be used to photograph the screen with the exposure time being equal to that required for at least one complete scanning raster.
  • Suitable additional components for the above-described control circuit 16 by which various electronic image enhancement techniques may be utilized if desired and by which radiographic image data may be stored on magnetic tape or the like instead of utilizing the cumbersome conventional film storage, are described in applicant's prior United States Patent No. 3,949,229 and may be embodied in the present system if desired.
  • the electrical control circuit 16 of this example will hereinafter be described in more detail in connection with certain additional provisions which may be employed in the circuit to alleviate forms of optical distortion in the image which can otherwise be present.
  • a focal length adjustment may be needed where a single instrument is to be used for the production of both wide-angle panoramic pantomographic images and also periapical dental radiographs or other images in which a smaller portion of the subject is to be imaged at a larger magnification. This accomplished by replacing the probe 14A with another probe which locates the detector 44 further out from the target anode plate 24 when focal length is to be increased or which locates the detector closer to the target anode plate when focal length is to be decreased. When the location of the detector 44. is changed in this manner, collimator 29A must also be replaced with another collimator having radiation transparent passages 68 convergent towards the new location of the detector 44.
  • Still another adjustment may also be made in conjunction with such a change of probe and collimator, In particular, if the region within the X-ray tube where the electron beam 23 is being deflected from the central axis of the tube is. spaced from the target anode plate 24 a distance substantially different from the spacing of the detector 44 from the target anode plate a distortion of the image can occur.
  • the axial position of the deflector 31 is also shifted by the previously described means so that the electron beam deflection region and the detector 44 remain substantially equidistant from, or in constant distance ratio to, the target anode plate surface 26.
  • Figure 5 depicts the forward portion of the X-ray tube 13 after adjustments have been made as discussed above in order to increase the focal length.
  • the original probe has been replaced with a new probe 14B which may have an internal construction essentially similar to that of the probe described above except that it is of greater length to position the detector 44B further out from the face of the tube.
  • the curvature of the probe 14B may also be modified as necessary to maximize patient comfort.
  • the original collimator has been replaced with a second collimator 29B in which the radiation-transmissive passages 68B have a different angular orientation in order to be convergent at the new, more distant, location of the detector 44B. It may be observed that collimator 29B also differs from that depicted in Figure 1 by being of the flat configuration previously described.
  • probe 14 described in connection with Figures 1 to 5 are primarily designed for production of pantomographic dental X-rays and have been shown and described as being mounted at one lateral side or the other of the face of the X-ray tube.
  • the probe may have a considerably different configuration and may be attached to the X-ray tube at a different location.
  • Figure 6 depicts the forward portion of the X-ray tube 13 as adjusted to provide for the making of periapical X-ray images of individual teeth or of a small number of individual teeth of the lower jaw of the patient.
  • a modified probe 14C is mounted on the face of the X-ray tube at the particular attachment means 52D which is situated above the central axis of the tube.
  • the modified probe 14C has a different configuration from the previously described probe and extends directly forward from the face of the tube and has a slight downward curvature towards the remote or distal end of the probe. In this example this situates the detector 44C at a differing distance from the face of the tube than in the case of the other probe and therefore a different collimator 29C is utilized which has a correspondingly changed focal length as hereinbefore described.
  • the same modified probe 14C may then be used to make periapical X-rays of teeth of the upper jaw by removing the probe from attachment means 52D and remounting it in the lower attachment means 52C below the collimator 29C as shown by dashed lines 14C' in Figure 6.
  • Use of the apparatus of Figure 6 to make a periapical X-ray image of the upper incisor teeth of a patient, with the probe at dashed line position 14C' is depicted in Figure 7.
  • the X-ray tube 13 may be positioned directly in front of the patient's upper jaw with the probe 14C' extending into the patient's mouth and slightly upwardly to locate the detector 44C of the probe directly behind the upper front incisor teeth of the subject 12.
  • a mounting bracket 88 may be secured to housing 27 at the underside of the housing in this instance for connection to suitable support means 89.
  • the support means 89 in this example is a semi-rigid tubular gooseneck of the form which can be bent to different configurations by applying sizeable force but which is otherwise sufficiently rigid to support the X-ray tube 13 in a selected position and orientation.
  • scissors brackets or the like of the known forms used to support more conventional dental X-ray tubes may be used.
  • Support means 89 may, if desired, attach to a housing 91 which contains electrical control circuit components of the system in which case conductors for coupling the electrical components of the tube with the rest of the circuit may be situated within the support tube 89.
  • FIG. 8 illustrates the forward portion of the X-ray tube 13 supporting still another probe 14D which is designed for facilitating the making of a medical X-ray image of the central region of a patient's head and showing, among other anatomical features, the bony structure of the ear.
  • the probe 14B may be mounted in one of the side attachment means at the face of the tube, attachment means 528 in this case.
  • Probe 14D is of greater length than those previously depicted and described and is shaped to curve around the head of the patient and to enter a small distance into the ear canal at the side of the patient's head opposite from the face of the X-ray tube.
  • probes having still other configurations may be provided to locate the detector within cavities of manufactured industrial parts, such as castings, for example, which are to be inspected by X-ray imaging.
  • a first method is to increase or decrease the size of the scanning raster at the visual display device 17 relative to the scanning raster size at the X-ray tube 13.
  • the other method is to change the position of the subject relative to the target anode plate 24 and detector 44. If the subject is positioned relatively close to the detector 44 it is highly magnified in the image although the field of view is reduced. Conversely, if the subject is repositioned closer to the target anode plate 24 and further from the detector 44, magnification is decreased but a wider-angle view is obtained.
  • the first method of magnification control can readily be effected simply by adjusting the beam deflection controls of either the X-ray tube 13 or the visual display device 17 or both.
  • the second method is accomplished by simply moving the X-ray tube 13 and probe 14 relative to the subject or vice versa.
  • a second factor which can cause variable magnification at different areas of the image arise from the fact that if the rate of scanning of the electron beam along the target anode plate 24 is uniform, then the effective rate of scanning along the subject to be imaged may not be uniform.
  • the teeth of the dental arch 12 to be imaged in this example are situated approximately along a circular arc A having a radius R and having a center of curvature D at the detector 44. If the electron beam of X-ray tube 13 is scanned in the X-direction, that is in the plane of Figure 9, at a uniform speed, the effective scanning rate along arc A is variable.
  • Effective scanning rate is slowest along the central portions of arc A which are closest to being parallel to the target anode plate 24 but progressively increases toward the end portions of arc A which increasingly curve away from the plane of the target anode plate.
  • the practical effect is that the central teeth occupy a disproportionately large amount of the image in the X or lateral direction or, in other words, tend to be more greatly magnified in the X direction than are the teeth nearer the sides of the image.
  • variable magnification effects described above may be reduced or substantially eliminated by delinearizing the scanning sweep speeds in both the X and the Y directions in either the X-ray tube 13 or the visual display device 17 in order to compenstate for the effects described above. It is preferable to delinearize the scanning action of the X-ray tube 13 for this purpose, and electrical circuits for this purpose will be hereinafter described. In order to best understand the operation of such circuits, a more mathematical analysis of the above-described effects is required.
  • point 0 designates the center of the target anode plate 24 which is the point of impact of the undeflected electron beam along the axis of the tube and defines the origin point of the coordinate system for the following equations.
  • Point P represents an arbitrarily chosen point of electron beam impact on the target anode plate in the course of a scanning raster.
  • the letter D designates the position of the detector 44 at the center of curvature of arc A along which the teeth to be imaged are situated while R is the radius of the arc.
  • C designates the fixed distance of arc A from point 0 along the central ray path 0-D while V is a variable representing the distance of arc A from the momentary electron beam impact point P.
  • X is a variable representing the displacement of point P from point 0 in the X-scan direction and 0 is a variable representing the angle formed by ray lines O-D and P-D.
  • S is the coordinate distance of ray line P-D from ray line 0-D measured along arc A.
  • Figure 10 is a diagrammatic section view taken along line X-X of Figure 9 or in other words along a vertical plane containing the central ray path O-D.
  • the tooth T1 which that plane intercepts requires a Y-direction sweep distance of Y o in order to be fully imaged and is magnified by the factor (R+C)/R for the reasons hereinbefore discussed.
  • Figure 11 is a diagrammatic section view taken along line XI-XI of Figure 9 or in other words along a vertical plane containing the ray path P-D.
  • the different tooth T2 which the plane of Figure 11 intercepts requires a Y-direction sweep distance of Yp in order to be fully imaged and is magnified by the factor (V+R)/(C+R) relative to the magnification of tooth T1.
  • tooth T1 and T2 of Figures 10 and 11 respectively are actually of the same height, tooth T2 is more greatly magnified in the image than is tooth T1.
  • the objective of the correction is to cause all teeth in the image to exhibit the same height Y o since as postulated for the present analysis, that is the actual fact. Mathematically this may be derived as follows:
  • F(t) is a triangular wave function owing to the lack of significant curvature of the teeth in the vertical direction.
  • the Y-direction sweep rate of the electron beam in the X-ray tube 13 should be deiinearized and caused to vary as a secant function of 0, circuit means for this purpose being hereinafter described.
  • secant function for this purpose results from the disposition of the teeth along an arc A having a constant radius of curvature R. If the object or series of objects to be imaged lie along a path having some other configuration wherein the radius of curvature is not constant, but instead varies as a function of 0, then a more complicated function than the secant function must be generated by essentially similar trigonometric analysis.
  • the teeth which are to be imaged in this example are of different thickness along the X-ray paths from the target anode plate 24 to detector 44.
  • the X-ray energy level best suited for producing a clear image of one tooth may not be the value best suited for imaging others of the teeth which have a different radiolucence.
  • the teeth tend to increase in thickness and therefore in radio-opacity from the front of the patient's dental arch 12 towards the back or, in the depicted geometry, in the direction of increasing 0.
  • the X-ray absorbency of the subject to be imaged is lowest at the extreme minus 0 portion of the X-direction scan and tends to increase as the angle 0 approaches its maximum positive value.
  • Clarity of the image may be enhanced by changing the energy of the electron beam within the X-ray tube 13 in the course of the scanning action in a programmed manner which compensates for the variations of radiolucence of different regions of the subject which have been described above, circuit means for this purpose being hereinafter described.
  • this may take the form of progressively increasing electron beam energy, in the course of each X-scan, as a function of 0.
  • the energy change in the course of each scan need not be a continuous gradual rise but may consist of a series of stepped increases or decreases each determined by the radiolucence of the particular tooth or portion of a tooth being imaged at successive stages in the scan.
  • an acceptable degree of compensation may often be achieved by simply changing electron beam energy level once in the course of each scan at the stage where the scan passes between the relatively thick back or molar teeth and the more radiation-transparent anterior teeth.
  • a first such effect is the attenuation of a radiation flux with distance in accordance with the well-known inverse square law.
  • ray path P-D is longer than the central ray path 0-D and in general the distance which X-rays must travel from the target anode plate 24 to detector 44 progressively increases from the center of the X-scan towards each extremity of such scan. This is also true of scanning action in the Y-direction. Owing to inverse square law attenuation this will cause count rate variations at the detector 44 even if a subject of uniform X-ray opacity is being imaged or if there is no X-ray absorber at all between the X-ray source and the detector.
  • the target anode plate 24 at which X-rays originate has a finite thickness which is not evident in Figure 9 but which may be seen by reference to Figure 12 wherein such thickness, designated T, has been greatly exaggerated for clarity of illustration.
  • An arbitrary point of electron impact on target anode plate 24 is identified by the letter E in Figure 12 and its coordinates are shown as r, 0 and 0 with such coordinates and other symbols having the same meaning as in the previous figures and discussion except that it should be understood that r, the distance from the X-ray origin point to detector 44, includes the thickness T of the target anode plate.
  • X-rays originating at point 0 at the center of the scanning raster pass directly through the target anode plate. 24 to reach the detector 44 and are attenuated to some specific degree because of absorption by the target anode plate material.
  • X-rays originating at an off-center point such as E must pass obliquely through the thickness T of the target anode plate and are therefore attenuated to a greater degree.
  • the effective length, of the X-ray path through the target anode plate, designated T" progressively increases towards the extremities of the scan in both the X and Y directions. Thus, a corresponding unwanted variation of X-ray count rate occurs at detector 44 unless a correction is provided.
  • Correction for both the inverse square law form of X-ray count variation and the variable effective target thickness form of X-ray count variation may be realized by varying the electron beam current within the X-ray tube 13 during the scanning operations to adjust radiation flux in such a manner as to compensate for these effects.
  • X-ray count rate at the detector 44 varies as a function of the ratio (R+C) 2 /r 2 .
  • electron beam generating and controlling elements of the X-ray tube 13 include a cathode 101 which emits electrons upon being heated by a filament 102 in the conventional manner and which has a terminal 101' to which a high negative electrical potential is applied from a high-voltage supply 103 in order to produce an electrical field which accelerates the electrons towards the grounded target anode plate 24 of the X-ray tube.
  • Maintaining the cathode 101 at a high voltage while grounding the anode 24 is the reverse of the conventional arrangement in the X-ray tubes and offers the advantage that a dental or medical patient or an electrically conductive inanimate object may be placed very close to the face of the X-ray tube or even against the face of the X-ray tube without creating a risk of electrical shock. Where this is not a problem, the cathode may be grounded and a positive high voltage supply may be connected to the anode plate, if desired.
  • the high voltage supply 103 is of the programmable form in which the magnitude of the output voltage delivered to cathode 101 is adjustable by varying the input voltage signal so that the electron beam energy may be modulated in the course of the scanning action as will hereinafter be discussed in more detail.
  • the electron gun 22 of X-ray tube 13 also has a control grid 105 with a terminal 105' to which a voltage may be applied for the purpose of modulating electron beam current as will also be discussed in more detail.
  • the electron gun 22 may also have further elements such as a first anode 104 and ultorfocusing grid 106 which are not utilized in accomplishing the image corrections of the present invention but which are present for their conventional purposes.
  • the X-ray tube 13 also has an X-beam deflection coil 33X and a Y-beam deflection coil 33Y for controlling the point of impact of the electron beam on target anode plate 24, the coils respectively having an X-deflection signal terminal 78 and a Y-deflection signal terminal 82.
  • the control circuit is further provided with a regulated DC power supply 107 having a first output terminal B+ at which a constant positive voltage is provided for operating other components of the circuit and a second output terminal B- at which a constant negative voltage is provided for similar purposes.
  • a regulated DC power supply 107 having a first output terminal B+ at which a constant positive voltage is provided for operating other components of the circuit and a second output terminal B- at which a constant negative voltage is provided for similar purposes.
  • the power supply connections to most components of the circuit are not depicted where such connections may be of the conventional known form.
  • the X-sweep frequency generator 77 has an output terminal 77' connected to the X-sweep frequency terminal 79 of the visual display device 17 and may be of the known construction which generates a ramp signal output wave form of the type depicted at 77W which cyclically oscillates in a linear manner between a maximum negative voltage and a maximum positive voltage at a frequency corresponding to the desired X-sweep or scan frequency at both the X-ray tube 13 and the visual display device 17.
  • the Y-sweep frequency generator 81 has an output terminal 81' coupled to the Y-sweep frequency terminal of visual display device 17 and is of the known form that produces a ramp signal 81W similar in general form to that of the X-sweep frequency generator except insofar as it has a substantially lower frequency as determined by the number of horizontal scan lines which are desired in the scanning raster.
  • the difference between the X and Y-sweep frequencies is normally much greater than appears from the wave forms 77W and 81 W in Figure 13, it being impractical to illustrate the actual difference because of space limitations in the drawings.
  • Optical X-ray count signals originating at the detector 44 are converted to electrical signals by photomultiplier tube 46 as previously described.
  • the X-ray count signal output terminal 84 of photomultiplier tube 46 is coupled to the Z or intensity signal terminal 86 of the visual display device 17 through an adjustable gain amplifier 108 and then through one input of a differential amplifier 109.
  • the other input of the differential amplifier 109 is connected to a selectable DC voltage source 111 which may consist of an adjustable contact 112 movable along a resistive element 113 having opposite ends connecting to the B+ and B- terminals of the DC power supply 107.
  • Adjustable gain amplifier 108 enables selective control of the voltage level of the intensity signal applied to the visual display device while the differential amplifier 109 and selectable voltage source 111 aids in suppressing detector noise by suppressing electrical pulses of less than a selected amplitude.
  • This form of intensity signal channel is best adapted to instances where the average rate of X-ray counts at detector 44 is sufficiently low that individual counts are separated in time and can be distinguished and individually processed.
  • an integrating form of Z signal channel may be utilized as described for example in prior United States Patent No. 3,949,229.
  • control circuit as described to this point can be utilized without further complication to produce an image at display device 17 if output terminal 77' of the X-sweep frequency generator is connected to X-deflection coil terminal 78 of the X-ray tube, output terminal 81' of the Y-sweep frequency generator is connected to terminal 82 of the X-ray tube and constant voltages are applied to the cathode terminal 101' and control grid terminal 105' of the X-ray tube but in that event each of the previously described forms of image distortion and image clarity degradation may be present.
  • Such images may be useful under many circumstances' particularly where the detector 44 is spaced a substantial distance away from the target anode plate 24 of the X-ray tube since the severity of several of these effects is an inverse function of the focal length of the system as determined by the degree of convergence of the various X-ray paths from the tube toward the detector.
  • compensation circuit means 114 are connected between the sweep frequency generators 77 and 81 and the X-ray tube 13 terminals in order to modify the X- and Y-sweep frequency wave forms and to modulate the electron beam energy and current in accordance with the several controlling mathematical relationships hereinbefore derived.
  • Compensation circuit means 114 relies primarily on a series of electronic function generators 116, 117, 118, 119 and 121 and multipliers 122, 123 and 124 which are depicted in block form in Figure 13 inasmuch as such circuit components may be of known internal construction and are available commercially.
  • function generators of this kind produce an output voltage which varies, as a function of an input voltage, in accordance with a predetermined mathematical relationship.
  • Function generator 116 for example is of the known form which produces an output voltage that varies as a tangent function of an input voltage while function generator 117 is of the form producing an output voltage that varies as a secant function of the input voltage.
  • Function generators 118 and 119 are of the form which produces an output voltage proportional to the square of the input voltage and function generator 121 produces an output voltage proportional to the square root of an input voltage.
  • Multipliers 122, 123 and 14 are each of the known form that produce an output voltage proportional to the product of two voltages applied to two inputs of the multiplier. Function generators and multipliers of this general form are often used for example in analog computer circuits. While the design of such function generators is known, a suitable internal circuit for a representative one of the function generators 116 will be hereinafter described in order to facilitate an understanding of certain characteristics of the compensation circuit means 114 as a whole.
  • variable magnification effects in the X-scan direction may be compensated for by varying the X-sweep frequency signal as applied to terminal -78- of the X-ray tube 13 as a function of the tangent of 0 rather than in the linear manner produced by the X-sweep frequency generator 77.
  • output terminal 77' of the X-sweep frequency generator may be coupled to the input of tangent function generator 116 through an adjustable gain amplifier 141 N which enables adjustment of signal level to determine the length of the X-direction scan in the X-ray tube.
  • the magnetic deflection system employed in the present example of the X-ray tube 13 is responsive to current through the X-deflection coil 33X, rather than to voltage as such as in the case of an electrostatic deflector, the output of tangent function generator 116 is coupJed to X-ray tube terminal 78 through a current amplifier 127 and variable resistor 128. If an electrostatic deflection system is used in X-ray tube 13, a voltage amplifier may be substituted for current ampli- _ tier 127.
  • the input X-sweep frequency wave form 77W applied to the tangent function generator 116 input terminal .116i may be treated as consisting of a sequence of triangular electrical pulses of positive polarity alternated with similar but inverted pulses of negative polarity.
  • the upper half of the circuit of Figure 14 processes the positive portions of the input wave form 77W while the lower half of the circuit processes the negative portions of the input wave form.
  • the input wave form 77W is applied to one input of a differential amplifier 129P which has a reference input connected to a selectable DC voltage source 131 P which may be adjusted to cause amplifier 129P to suppress the negative portions of the incoming wave form while transmitting the positive portion on as a series of triangular positive pulses separated in time.
  • a first diode 132P and variable resistor 133P are connected between amplifier 129P and a summing junction 134.
  • a second diode 136P and variable resistor 137P are also connected between amplifier 129P and summing junction 134 except that in this instance the input of diode 136P connects to the output of amplifier 129P through the adjustable contact of a potentiometer 139P having a resistive element connected between the output of amplifier 129P and the negative power supply terminal B-.
  • a third diode 141 and variable resistor 142P are connected between the output of amplifier 129P and the summing junction 134 through the movable contact of another potentiometer 143P having a resistive element connected between the output of amplifier 129P and power supply terminal B-.
  • Summing junction 134 is in turn coupled to the output terminal 116 0 of the tangent function generator through an operational amplifier 144 and a resistor 146 connected in parallel with the operational amplifier.
  • Figure 15A illustrates the wave form 77W of the X-sweep frequency signal applied to the input 116i of the tangent function generator of Figure 14.
  • Figure 15B illustrates the modified output wave form of the tangent function generator at output terminal 116 0 .
  • the output voltage rises in a non-linear manner in a series of three voltage rise segments, a, b and c each of which taken individually is linear but which are of progressively increasing slope.
  • the positive output voltage then decreases in a similar series of linear segments of progressively diminishing slope.
  • the circuit of Figure 14 modified the triangular input wave form in this manner since as the leading edge of the positive triangular pulse appears at the output of amplifier 129P, diode 132P conducts immediately to transmit a rising current to summing junction 134. Immediate conduction of diode 132P is provided for by adjusting selectable voltage source 131 P to offset the forward bias of the diode with a base voltage output from amplifier 129P. This corresponds to the initial segment a of the wave form of Figure 15B. Diodes 136P and 146P do not initially conduct because of the respectively more negative biases applied to the inputs of such diodes by potentiometers 139P and 143P respectively.
  • the bottom half of the circuit of Figure 14 is essentially similar to the top half except insofar as the diodes 132N, 136N and 141 are inverted relative to the counterpart diodes of the top half of the circuit and except insofar as the potentiometers 131 N, 139N and 143N are connected to the B+ power supply terminal rather than the B- terminal as in the case of the counterpart potentiometers of the upper half of the circuit.
  • the lower half of the circuit modifies the negative portions of the incoming wave form in essentially the same manner that the upper half of the circuit modifies the positive portions of the incoming wave form and the two portions as modified are combined at summing junction 134 to produce the complete modified output wave form 77T at the output terminal 116o.
  • this form of function generator does not produce the desired wave form in an idealized form free of discontinuities but instead approximates the desired wave form by a series of segments a, b and c which in themselves are linear.
  • the remaining distortion in the image arising from this departure from an ideal modified sweep signal is sufficiently small that it does not present any practical problems is most cases.
  • the circuit of Figure 14 may be modified by adding additional stages of the diodes 132, 136, 141, variable resistors 133, 137, 142 and potentiometers 139 and 143 on both the positive upper side and the negative lower side of the circuit so that the output wave form as shown in Figure 15B is modified by having a greater number of the linear segments a, b and c, each of shorter duration than in the present case, so that the desired idealized wave form is even more closely approached.
  • output 81' of the Y-sweep frequency generator is coupled to one input of multiplier 122.
  • Output 77' of the X-sweep frequency generator is coupled to the input of function generator 117 through an adjustable gain amplifier 147.
  • Function generator 117 produces an output wave form representing the secant function of 0 and such output is transmitted to the other input of multiplier 122 through another adjustable gain amplifier 148.
  • the output of multiplier 122 is a voltage wave form 122W corresponding to the linear Y-sweep frequency generator output signal 81 W as delinearized to vary as a secant function of angle 0 during each successive X-direction scan.
  • the output of the multiplier 122 is typically a low current voltage signal whereas the magnetic Y-deflection coil of the X-ray tube requires a relatively high current
  • the output of multiplier 122 is couled to the Y-deflection signal terminal 82 of the X-ray tube through an adjustable gain amplifier 149 and a current amplifier 151 and variable resistor 152.
  • still another function generator 153 may have an input connected to the X-sweep frequency output terminal 77' through an adjustable gain amplifier 154 and has an output connected to the voltage control signal terminal 103' of programmable high voltage supply 103 through another adjustable gain amplifier 156. If the variation of X-ray energy in the course of the X-scan is always to be the same for all usages of the X-ray tube 13, then the function generator 153 may be of the form which produces some single predetermined modification of the input signal with corresponds to the desired change of beam energy in the course of the scanning action.
  • the function generator 153 may be of the fixed form which simply produces a gradual relative rise of the output signal as compared to the input signal in order to cause the programmable high voltage supply 103 to progressively increase electron beam energy during the course of each X-scan line of the X-ray tube 13.
  • function generator 153 is of the form which enables selection of any of a variety of functions in order to readily accommodate to different usages of the X-ray tube 13 and to subjects which may have differing patterns of varying radiolucence in the X-direction.
  • the function generator circuit of Figure 14, for example, offers considerable latitude for selection of different output waveforms by selected adjustments of the several variable resistances 133, 137, 142 and potentiometers 131, 139 and 143.
  • Other forms of function generator enabling an even greater variety of predetermined wave fbrm modulations are known to the art and may be utilized if desired.
  • the remaining components of the circuit means 114 compensate for the previously described unwanted variations of radiation flux level at the subject and at the detector arising from inverse square law effects and the variation of absorption of X-rays in the target anode plate 24 at different portions of the scanning action.
  • the previously derived mathematical expression (17) is the controlling relationship in accordance with which the voltage applied to the control grid 105 of the X-ray tube must be varied in the course of the scanning action in order to vary electron beam current in such a manner as to compensate fbr these effects.
  • expression (17) a considerably more complex correction function is involved than is the case with the other forms of correction described above.
  • multiplier 123 receives the output of secant 0 function generator 117 through squaring function generator 118.
  • the one input of multiplier 123 receives a sec2( ⁇ signal.
  • Output 81' of the Y-sweep frequency generator is coupled to the input of function generator 119, which is also a squaring module of the form that produces an output voltage proportional to the square of the input voltage.
  • the output of squaring function generator 119 is connected to one input of a differential amplifier 157 having an output connected to the remaining input of multiplier 123.
  • the other input of differential amplifier 157 is connected to a selectable voltage source 158 which may be of the form having a resistive element connected across the B+ and B- terminals of the power supply and having a movable element connected to the reference input of differential amplifier 157.
  • Selectable voltage source 158 is adjusted to generate a voltage representing the constant (R+C) 2 term of mathematical expression (13).
  • the output of squaring function generator 119 is a representative of the term Yo of that same mathematical expression and the two are summed by amplifier 157
  • the output of amplifier 157 is the r o 2 term of the controlling mathematical expression (17). This r o 2 signal voltage is multiplied by the se C 2 0 signal voltage from function generator 118 in multiplier 123.
  • the output voltage from multiplier 123 is transmitted through a selectable gain amplifier 158 which is adjusted to scale the input signal by the constant (uT/R+C) 2 term of mathematical expression (17) and the output of amplifier 158 is then passed through the square root function generator 121 to produce a voltage signal representing the function (uT/R+C) (r o sec 0) at one input of multiplier 124.
  • the other input of multiplier 124 receives the r o 2 sec 2 ⁇ voltage signal from the output of multiplier 123.
  • the output of multiplier 124 represents (uT ro3sec3( ⁇ /(R+C) which is the second term in the sum given by mathematical expression (17).
  • the correction factor K of expression (17) is then obtained by adding in the expression r o 2 sec 2 ⁇ which as previously described is available at the output of multiplier 123.
  • an adding amplifier 159 has one input connected to the output of multiplier 124 and has the other input connected to the output of multiplier 123. This results in an analog voltage appearing at the output of amplifier 159 which is representative of the desired correction factor K as defined in the previously given mathematical expression (17).
  • the output of amplifier 159 is connected to the control grid terminal 105' of the X-ray tube in order to modulate electron beam current to provide the desired image correction.
  • adjustment of the several adjustable controls of the circuit of Figure 13 to minimize image distortion may be facilitated by using a coarse, preferably flexible gauze material formed of X-ray opaque wires or the like.
  • a gauze may be wrapped around the subject to be imaged and the scanning X-ray system may then be temporarily operated at a low power level while viewing the image of the gauze on the screen of the visual display device.
  • the several controls of the distortion correction circuit then readily be adjusted to bring the images of the wires of the gauze into parallelism in both the horizontal and vertical directions with the intervening open spaces being equalized. After linearization of the low-level image has been achieved, the gauze may be removed and the electron beam energy of the tube increased to the normal operating level for obtaining the desired radiograph.
  • circuit of Figure 13 includes components for providing each of the several forms of image correction hereinbefore discussed, there are instances where at least some of the forms of distortion or image degradation are not significant enough to need correction in which case the corresponding corrective component portions of the circuit of Figure 13 may be omitted.
  • the small X-ray detector of the previously described embodiments may simply be positioned further out from the face of the X-ray tube and a modified collimator may be used which has radiation-transmissive passages which are less convergent and which are directed at the more distant X-ray detector.
  • a modified collimator may be used which has radiation-transmissive passages which are less convergent and which are directed at the more distant X-ray detector.
  • such an arrangement produces other adverse effects which outweigh the advantages of lessened distortion.
  • the image would include the particular teeth which are intended to be in the image but superimposed thereon would also be images of teeth or other anatomical structures at the other side of the dental patient's head and such images would tend to obscure the desired images of the particular teeth of interest.
  • the embodiment of Figure 16 utilizes a modified detector 44E which differs from that of the previously described embodiments by being larger and by being situated between the tube 13E and the focal point 161 rather than at the focal point.
  • the radiation-sensitive area of the detector 44E is preferably just large enough to intercept the X-rays travelling from the scanning raster at target anode plate 24E towards the focal point 161 since the configuration of the subject may limit detector size. In other words, a very large detector cannot be inserted into constricted spaces such as into the mouth of the dental patient of Figure 16.
  • the embodiment of Figure 16 is particularly appropriate for periapical rather than panoramic imaging.
  • the smaller field of view which may be in some cases dictated by the restriction on detector 44E size, is characteristic of periapical radiography.
  • the embodiment of Figure 16 may employ an X-ray source 13E of smaller size having a smaller scanning raster area.
  • panoramic images may still be obtained by shifting the X-ray tube and detector and making a series of images.
  • the detector 44E and X-ray tube 13E may be shifted to the dashed-line positions 44E' and 13E' for example by rotation along an arc 163 having a rotational axis at the approximate center of curvature of the portion of the dental arch 12 being imaged.
  • the X-ray tube 13E and detector 44E may also be translated in a direction parallel to that rotational axis so that both maxillary and mandibulary teeth or other cranial features may be imaged. If desired, all of the images taken from these different viewpoints may be displayed on a single screen for viewing or to make up a mosaic of views combining the entire dental arch on one picture.
  • the camera shutter may be held open during the imaging process except during those periods when the X-ray tube 13E and detector 44E are in the process of being repositioned.
  • the detector 44E may, if desired, be supported and positioned by a probe 14E similar to those previously described and other portions of the system of Figure 16 may be similar to the corresponding parts of the previously described embodiments.
  • the additional collimator 164 is secured directly against the detector 44E as this maximizes effectiveness, minimizes collimator size and facilitates positioning and support of the collimator. Secondary or scattered X-rays which are travelling in the general direction of the detector but at angles other than that of the X-rays from the primary collimator 29E are absorbed by the additional collimator 164 and thus do not produce erroneous signals in the detector. Under some circumstances, notably where the subject to be imaged is not a living organism for which radiation dosage should be minimized, the additional collimator 164 may be used without the primary collimator 29E.
  • Either collimator alone is capable of limiting X-ray transmission from the tube to the detector to a single path at a given instance in order to enable the production of an image.
  • the presence of a collimator between the subject and the X-ray tube is beneficial in that it suppresses X-rays which are not directed towards the subject along the particular path from which meaningful image data can be obtained.

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Claims (13)

1. Röntgenäbtastgerät (11) Erzeugung von Röntgenbilddaten, mit einem Gehäuse (27) in welchem eine Röntgenröhre (13) mit einer Vakuumhülle (19) einen Vakuumbereich (21) bildet, um eine Auffanganodenplatte (24) und einen Elektronenstrahlerzeuger (22) für die Erzeugung eines Elektronenstrahles (23) zu umschliessen, wobei die Röntgenstrahlröhre (13) ausserdem Elektronenstrahlablenkmittel (31) aufweist, um den Elektronenstrahl (23) zu aufeinanderfolgenden unterschiedlichen Punkten auf der Oberfläche der Auffanganodenplatte (24) zu richten, um so einen beweglichen Röntgenstrahaetiausgangspunkt auf der Auffanganodenplatte (24) zu erzeugen und wobei der Elektronenstrahlröhre (13) ein Sondenelement (14) zugeordnet ist, welches einen Röntgenstrahldetektor (44) trägt mit einem aktiven Bereich, welcher kleiner ist als der Bewegungsbereich des beweglichen Röntgenstrahlenausgangspunktes, um so die Röntgenstrahlen festzustellen, welche durch einen Gegenstand (12) hindurch gehen, welcher zwischen der Auffanganodenplatte (24) und dem Röntgenstrahlendetektor (44) angeordnet ist, wobei der Basisteil (51) des Sondenelementes (14) entfernt vom Röntgenstrahlendetektor (44) angeordnet ist, das Sondenelement (14) Mittel (48) hat, um Röntgenstrahlenzähldaten vom Detektor (44) zum Basisteil (51) des Sondenelementes (14) zu übertragen, soviel Aufnahmemittel (46,61) mit mindestens einer Ausgangsklemme (84) zur Übertragung elektrischer Signale, welche die Röntgenstrahlenzähldaten darstellen und eine Kollimatorvorrichtung (29), welche neben der Auffanganodenplatte (24) angeordnet ist, um Röntgenstrahlen von dem beweglichen Röntgenstrahlenausgangspunkt auf der Auffanganodenplatte (24) durch einen Gegensand (12) hindurch zu schicken zu dem Röntgenstrahldetektor (44), wobei die Kollimatorvorrichtung (29) darin gebildete Durchgänge aufweist, welche zum Detektor (44) konvergieren, dadurch gekennzeichnet, dass vorgesehen sind: (i) mehrere starre Sondenelemente (14A, 14B, 14C, 14D, 14E), welche entweder unterschiedliche Längen haben, um vorwählbar verschiedene feste Abstände zwischen dem Detektor (44) und der Auffanganodenplatte (24) festzulegen, oder unterschiedliche geometrische Formen haben, (ii) mehrere Kollimatoreinrichtungen (29A, 29B, 29C), mit unterschiedlichen Brennweiten, welche den festen Abständen entsprechen, die von dem Sondenelementen festgelegt sind, (iii) lösbare Befestigungsmittel (52) an dem Gehäuse zum Tragen eines der Sondenelemente derart, dass der Detektor (44) von ihnen in Abstand vor der Auffanganodenplatte (24) getragen ist, und (iv) weitere lösbare Befestigungsmittel (28, 76) um eine der Kollimatorvorrichtungen neben der Auffanganodenplatte (24) zu tragen, wobei die Anordnung derart ist, dass bei Anordnung einer ausgewählten Kollimatorvorrichtung in den zusätzlichen lösbaren Befestigungsmitteln (28, 76) ein entsprechendes Sondenelement ausgewählt und in den lösbaren Befestigungsmitteln (52) montiert werden kann, um einen Detektor (44) automatische im Brennpunkt der Kollimatoreinrichtung anzuordnen.
2. Röntgenabtastvorrichtung nach Anspruch 1, dadurch gekennzeichnet, dass eine weitere Kollimatorvorrichtung (29E) vorgesehen ist mit einer Brennweite, welche grösser ist als der Abstand zu einem Detektor (44E), welcher von einem der Sondenelemente getragen wird.
3. Röntgenabtastvorrichtung nach Anspruch 1 oder 2, dadurch gekennzeichnet, dass die Befestigungsmittel (52) für eine der Sondenelemente (14A, 14B, 14C, 14D, 14E) Mittel (48) umfasst, um optische Signale vom Detektorteil der Sonde zu einer lichtempfindlichen Vorrichtung (46) zu leiten, welche zur Erzeugung einer elektrischen Spannung in Abhängigkeit von Energiesignalen mit optischer Frequenz ausgelegt ist, sowie ein Übertragungselement (61) für optische Energie, welches sich von der lichtempfindlichen Vorrichtung (46) zu den Befestigungsmittel (52) erstreckt, dass der Röntgenstrahlendetektor (44) eine Szintillationsdetektor ist, und dass die Mittel (48) zum Weiterleiten der Röntgenstrahlenzähldaten von dem Detektor (48) zum Basisteil (51) des Sondenelementes (14) ein Lichtleiter ist, welcher in dem Sondenelement (14) enthalten ist.
4. Abtaströntgengerät nach Anspruch 3, dadurch gekennzeichnet, dass die Elektronenstrahlablenkvorrichtung (31) der Röntgenröhre (13) Mittel (41, 42) zur selektiven Änderung des Abstandes der Strahlablenkvorrichtung (31) bezüglich der Auffanganodenplatte (24) aufweist zur Anpassung an Änderungen des vorbestimmten Abstandes zwischen dem Röntgenstrahlendetektor (44) und der Auffanganodenplatte (26), welche durch Änderungen der Sondenelemente (14A, 14B, 14C, 14D, 14E), hervorgerufen werden.
5. Röntgenabtastvorrichtung nach einem der Ansprüche 1-4, gekennzeichnet, durch eine Platte aus Strahlenfiltermaterial (73), welche neben dem Strahlenkollimator (29) und im Weg der Röntgenstrahlen, welche durch den Strahlungskollimator (29) hindurchgewandert sind, angeordnet ist.
6. Röntgenabtastvorrichtung nach einem der Ansprüche 1-5, dadurch gekennzeichnet dass der Strahlungskollimator (29) einen Block aus strahlungsabsorbierendem Glas (67) ist, welcher von einer Vielzahl von einander getrennten strahlungsübertragenden Durchgängen (68) durchbohrt ist, wobei jeder der Vielzahl von strahlungsübertragenden Durchgängen (68) auf einen einzigen Punkt gerichtet ist, welcher sich nach aussen hin in einem bestimmten Abstand von der Oberfläche des Blocks strahlungsabsorbierenden Glases (67) befindet.
7. Röntgenabtastvorrichtung nach Anspruch 6, dadurch gekennzeichnet, dass der Block strahlungsabsorbierenden Glases (67) eine Vielzahl rohrförmiger optischer Glasfaserelemente umfasst, welche miteinander verschmolzen sind um den Block strahlungsabsorbierenden glases (67) zu bilden durch den die strahlungsübertragenden Durchgänge (68) hindurchgehen.
8. Röntgenabtastvorrichtung nach einem der Ansprüche 1-7, dadurch gekenzeichnet, dass die Elektronenstrahlablenkvorrichtung (31 mit der der Elektronenstrahl auf eine Fläche der Auffanganodenplatte (24) der Röntgenstrahlröhre (13) gerichtet wird, ausgelegt ist, um den Elektronenstrahl wiederholt über einen Rastermusterbereich auf der Oberfläche in einer X-Koordinaten-Richtung zu schwenken und die Lage der X-Koordinaten-Richtung wiederholt in Richtung einer Y-koordinaten-Richtung zu versetzen.
9. Röntgenabtastvorrichtung nach Anspruch 8, dadurch gekennzeichnet, dass die Elektronenstrahlablenkvorrichtung (31) von einem Steuerkreis (16) gesteuert wird, welcher einen X-Ablenkfrequenzsignalgenerator (77) zur Erzeugung eines elektrischen X-Ablenkfrequenzsignals umfasst, um eine Signalablenkung in X-Koordinaten-Richtung zu erzeugen, sowieeinen Y-Ablenkfrequenzgenerator (81) zur Erzeugung eines Y-Ablenkfrequenzsignals zur Erzeugung der Versetzung in Y-Koordinaten-Richtung und einen Kreis zur Übertragung des X-Ablenkfrequenzsignals vom X-Ablenkfrequenzsignalgenerator (77) und das Y-Ablenkfrequenzsignals vom Y-Ablenkfrequenzsignalgenerator (81) zur Strahlablenkvorrichtung (31) der Röntgenstrahlröhre (13), wobei der Kreis dadurch gekennzeichnet ist, dass er einen ersten Funktions-generator (116) umfasst, werden zwischen den X-Ablenkfrequenzgenerator (77) und die Strahlablenkorrichtung geschaltet ist, um die Geschwindigkeit der Ablenkung des Elektronenstrahles an der Auffanganodenplatte in X-Koordinaten-Richtung zu entlinearisieren.
10. Röntgenabtastvorrichtung nach Anspruch 9, dadurch gekennzeichnet, dass der erste Funktionsgenerator (116) die Geschwindigkeit der Röntgenstrahlabtastung des Gegenstandes in X-Koordinaten-Richtung linearisiert durch Entlinearisierung der Geschwindigkeit des Abtastens des Elektronenstrahles an der Auffanganodenplatte in X-Koordinaten-Richtung.
11. Röntgenabtastvorrichtung nach Anspruch 9 oder 10, gekennzeichnet durch einen zweiten Funktionsgenerator (117, 122), welcher zwischen den X- und Y-Ablenkfrequenzgeneratoren (77, 81) und der Strahlablenkvorrichtung (31) angeordnet ist zur Linealisierung der Geschwindigkeit der Röntgenstrahlabtastung des Gegenstandes in Y-Koordinaten-Richtung durch Entlinearisierung der Geschwindigkeit der Abtastung des Elektronenstrahles an der Auffanganodenplatte (26) in Y-Koordinaten-Richtung.
12. Röntgenabtastvorrichtung nach einem der Ansprüche 9-11, dadurch gekennzeichnet, dass der Elektronenstrahlgenerator (22) der Röntgenstrahlröhre (13) ein Steuergitter (105) hat zur Veränderung des Elektronenstrahlstromes in Abhänghigkeit von Änderungen der anliegenden Steuergitterspannung und dass ein dritter Funktionsgenerator (117, 188, 190) zwischen die X-und Y-Ablenkfrequenzgeneratoren (77,81) und das Steuergitter (105) geschaltet ist, um die Steuergitterspannung im Verlauf der Ablenkung des Elektronenstrahles in X und Y-Richtung auf der Auffanganodenplatte in Abhängigkeit von Abstand des momentanen Auftreffpunktes des Elektronenstrahles auf der Auffanganodenplatte (24) von einem vorgegebenen Bezugspunkt in dem Abtastraster zu verändern.
EP78200018A 1977-06-03 1978-06-01 Röntgenabtastsystem und Verfahren. Expired EP0000079B1 (de)

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US05/803,077 US4196351A (en) 1977-06-03 1977-06-03 Scanning radiographic apparatus
US803077 1977-06-03

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EP0000079B1 true EP0000079B1 (de) 1983-01-12

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Families Citing this family (29)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US4239971A (en) * 1979-10-05 1980-12-16 Pennwalt Corporation Simultaneously displaying varying tomographic images of dental arch with single panoramic X-ray exposure
DE2949199A1 (de) * 1979-12-06 1981-06-11 Siemens AG, 1000 Berlin und 8000 München Zahnaerztliche roentgendiagnostikeinrichtung
US4519092A (en) * 1982-10-27 1985-05-21 Albert Richard D Scanning x-ray spectrometry method and apparatus
FR2547495B1 (fr) * 1983-06-16 1986-10-24 Mouyen Francis Appareil permettant d'obtenir une image radiologique dentaire
JPS60234645A (ja) * 1984-05-02 1985-11-21 フランシス ムヤン 歯科用放射線画像供給装置
US4730350A (en) * 1986-04-21 1988-03-08 Albert Richard D Method and apparatus for scanning X-ray tomography
US5237598A (en) * 1992-04-24 1993-08-17 Albert Richard D Multiple image scanning X-ray method and apparatus
US5267296A (en) * 1992-10-13 1993-11-30 Digiray Corporation Method and apparatus for digital control of scanning X-ray imaging systems
US5651047A (en) * 1993-01-25 1997-07-22 Cardiac Mariners, Incorporated Maneuverable and locateable catheters
US5682412A (en) * 1993-04-05 1997-10-28 Cardiac Mariners, Incorporated X-ray source
US5550378A (en) * 1993-04-05 1996-08-27 Cardiac Mariners, Incorporated X-ray detector
US5696806A (en) 1996-03-11 1997-12-09 Grodzins; Lee Tomographic method of x-ray imaging
US6183139B1 (en) 1998-10-06 2001-02-06 Cardiac Mariners, Inc. X-ray scanning method and apparatus
US6118854A (en) * 1998-10-06 2000-09-12 Cardiac Mariners, Inc. Method of making x-ray beam hardening filter and assembly
US6208709B1 (en) 1998-10-06 2001-03-27 Cardiac Mariners, Inc. Detection processing system
US6157703A (en) * 1998-10-06 2000-12-05 Cardiac Mariners, Inc. Beam hardening filter for x-ray source
US6421420B1 (en) 1998-12-01 2002-07-16 American Science & Engineering, Inc. Method and apparatus for generating sequential beams of penetrating radiation
DE19912854A1 (de) * 1999-03-22 2000-10-05 Sirona Dental Systems Gmbh Verfahren zur Korrektur des Vergrößerungsfaktors bei digitalen Röntgenaufnahmen
WO2000072354A1 (en) * 1999-05-25 2000-11-30 Dentsply International Inc. Dental x-ray apparatus
US7180981B2 (en) * 2002-04-08 2007-02-20 Nanodynamics-88, Inc. High quantum energy efficiency X-ray tube and targets
US7231017B2 (en) * 2005-07-27 2007-06-12 Physical Optics Corporation Lobster eye X-ray imaging system and method of fabrication thereof
US8842808B2 (en) 2006-08-11 2014-09-23 American Science And Engineering, Inc. Scatter attenuation tomography using a monochromatic radiation source
RU2428680C2 (ru) * 2006-08-23 2011-09-10 Эмерикэн Сайэнс Энд Энджиниэринг, Инк. Способ определения характеристик объекта
US7924979B2 (en) * 2006-08-23 2011-04-12 American Science And Engineering, Inc. Scatter attenuation tomography
US7714293B2 (en) * 2007-01-06 2010-05-11 General Electric Company Methods and apparatus for keystone effect
US20120305781A1 (en) * 2011-05-31 2012-12-06 Jansen Floribertus P M Heukensfeldt System and method for collimation in diagnostic imaging systems
EP2955510B1 (de) * 2011-11-25 2018-02-21 Aribex, Inc. Röntgenstrahlentfernungsanzeiger und zugehörige verfahren
US11191497B2 (en) * 2018-10-16 2021-12-07 Shayda Cullen Digital dental x-ray sensor device having a rounded housing including a radio transceiver
CN113433579B (zh) * 2021-05-18 2023-01-20 中国工程物理研究院激光聚变研究中心 一种大灵敏面x射线光谱平响应二极管探测器

Family Cites Families (12)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US2477307A (en) * 1946-11-09 1949-07-26 Mackta Leo Combined x-ray and fluoroscopic apparatus
US2638554A (en) * 1949-10-05 1953-05-12 Bartow Beacons Inc Directivity control of x-rays
US2730566A (en) * 1949-12-27 1956-01-10 Bartow Beacons Inc Method and apparatus for x-ray fluoroscopy
US2777068A (en) * 1952-08-27 1957-01-08 Joseph T Bowser X-ray film holders
GB1163107A (en) * 1965-12-01 1969-09-04 Siemens Ag Compensating for Vignetting in Television Systems.
JPS5318318B2 (de) * 1972-12-27 1978-06-14
US4032787A (en) * 1974-06-24 1977-06-28 Albert Richard D Method and apparatus producing plural images of different contrast range by x-ray scanning
US3949229A (en) * 1974-06-24 1976-04-06 Albert Richard D X-ray scanning method and apparatus
US4002917A (en) * 1974-08-28 1977-01-11 Emi Limited Sources of X-radiation
DE2506630A1 (de) * 1975-02-17 1976-08-26 Siemens Ag Zahnaerztliche roentgendiagnostikeinrichtung
US4031395A (en) * 1975-02-21 1977-06-21 Emi Limited Radiography
US4007375A (en) * 1975-07-14 1977-02-08 Albert Richard D Multi-target X-ray source

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US4196351A (en) 1980-04-01
JPS6233898B2 (de) 1987-07-23
JPS5416994A (en) 1979-02-07
EP0000079A1 (de) 1978-12-20
DE2862147D1 (en) 1983-02-17

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