Search Images Maps Play YouTube News Gmail Drive More »
Sign in
Screen reader users: click this link for accessible mode. Accessible mode has the same essential features but works better with your reader.

Patents

  1. Advanced Patent Search
Publication numberUS3854049 A
Publication typeGrant
Publication dateDec 10, 1974
Filing dateDec 10, 1973
Priority dateDec 10, 1973
Publication numberUS 3854049 A, US 3854049A, US-A-3854049, US3854049 A, US3854049A
InventorsF Kelcz, C Mistretta
Original AssigneeWisconsin Alumni Res Found
Export CitationBiBTeX, EndNote, RefMan
External Links: USPTO, USPTO Assignment, Espacenet
Compensation for patient thickness variations in differential x-ray transmission imaging
US 3854049 A
Abstract
Differential x-ray images can be produced by sub-tracting two different x-ray images which are produced by using two different quasi-monoenergetic x-ray spectra. Two different x-ray filters may be employed alternately to produce such spectra. For example, two different filters containing iodine and cerium may be employed alternately to produce two different x-ray spectra having peaks at different energies. The x-rays to be filtered may be derived from an ordinary x-ray tube which produces a continuous spectrum of x-rays over a wide band of energies. The purpose of producing the differential x-ray images is to subtract or cancel out the portions of the x-ray images which are due to the ordinary tissues of the patient, particularly the soft tissues, so that the presence of certain contrast substances in the patient's body will be emphasized or enhanced. Such contrast substances include iodine, xenon or barium, introduced into the bloodstream, the lungs or the food canal of the patient. The ordinary soft tissues of the patient transmit x-rays of different energies to different extents. While the two different x-ray images can be balanced for any particular thickness of the patient, variations in such thickness over the field of view tend to upset such balance. Thus, without compensation for patient thickness variations, cancellation of the x-ray images due to the ordinary tissues of the patient can be achieved at only one value of patient thickness. In accordance with the present invention, such patient thickness compensation is achieved by adjusting the composition and density of the two x-ray filters, and by adjusting the high voltage supplied to the x-ray tube, so that the ratio of the two different x-ray images produced by soft tissues remains nearly constant over a wide range of patient thickness. Thus, substantial cancellation of the soft tissue images can be achieved over a wide range of variations in the patient thickness. The x-ray filtration is adjusted so that one of the x-ray filters produces a spectrum of transmitted x-rays having two peaks at energies below and above the energy of the peak produced by the other x-ray filter. Thus, the sum of the images produced by the two peaks tends to remain in a constant relationship to the image produced by the other spectrum, despite variations in patient thickness. The supply voltage to the x-ray tube is varied as an inverse function of the average patient thickness. Thus, the voltage is reduced when the average patient thickness increases. The adjustment of the supply voltage to the x-ray tube changes the two different quasi-monoenergetic x-ray spectra so as to optimize the compensation for patient thickness variations.
Images(3)
Previous page
Next page
Description  (OCR text may contain errors)

United States Patent Mistretta et al.

[ Dec. 10, 1974 COMPENSATION FOR PATIENT THICKNESS VARIATIONS IN DIFFERENTIAL X-RAY TRANSMISSION IMAGING Primary Examiner.lames W. Lawrence Assistant Examiner-C. E. Church Attorney, Agent, or FirmBurmeister, Palmatier & Hamby [57] ABSTRACT Differential x-ray images can be produced by subtracting two different x-ray images which are produced by using two different quasi-monoenergetic x-ray spectra. Two different x-ray filters may be employed alternately to produce such spectra. For example, two different filters containing iodine and cerium may be employed alternately to produce two different x-ray spectra having peaks at different energies. The x-rays to be filtered may be derived from an ordinary x-ray tube which produces a continuous spectrum of x-rays over a wide band of energies. The purpose of producing the differential xray images is to subtract or cancel out the portions of the x-ray images which are due to the ordinary tissues of the patient, particularly the soft tissues, so that the presence of certain contrast substances in the patients body will be emphasized or enhanced. Such contrast substances include iodine, xenon or barium, introduced into the bloodstream, the lungs or the food canal of the patient. The ordinary soft tissues of the patient transmit x-rays of different energies to different extents. While the two different x-ray images can be balanced for any particular thickness of the patient, variations in such thickness over the field of view tend to upset such balance. Thus, without compensation for patient thickness variations, cancellation of the x-ray images due to the ordinary tissues of the patient can be achieved at only one value of patient thickness. In accordance with the present invention, such patient thickness compensation is achieved by adjusting the composition and density of the two x-ray filters, and by adjusting the high voltage supplied to the x-ray tube, so that the ratio of the two different x-ray images produced by soft tissues remains nearly constant over a wide range of patient thickness. Thus, substantial cancellation of the soft tissue images can be achieved over a wide range of variations in the patient thickness. The x-ray filtration is adjusted so that one of the x-ray filters produces a spectrum of transmitted x-rays having two peaks at energies below and above the energy of the peak produced by the other x-ray filter. Thus, the sum of the images produced by the two peaks tends to remain in a constant relationship to the image produced by the other spectrum, despite variations in patient thickness. The supply voltage to the x-ray tube is varied as an inverse function of the average patient thickness. Thus, the voltage is reduced when the average patient thickness increases. The adjustment of the 30 Claims, 12 Drawing Figures [MAGE IMAGE EE J SUBTRACTION TELEVISION 22/ svsrem S 72 25 MOVABLE 24 FILTRATION 16 x RAY SOURCE VARIABLE HIGH VOLTAGE l6 SUPPLY PATENIEL BEE I 01974 SHEET 10F 3 1D G0 .Z

IMAGE IMAGE C TV DE-QEDTOR j- SUBTRACTION 22/ SYSTEM svnc.

12 25 MOVABLE 24 FILTRATION 16 -x- RAY 24b SOURCE I Efi" X"RAY 2 3 ABSORPTION VARIABLE COEFFXCIENIS HIGH VOLTAGE l8 SUPPLY COMPENSATION FOR PATIENT THICKNESS VARIATIONS IN DIFFERENTIAL X-RAY TRANSMISSION IMAGING This invention relates to improvements in differential x-ray imaging techniques. A differential x-ray image can be produced by deriving the difference between two different x-ray images, produced under different conditions as to x-ray energy, or as to other factors. For example, the two different x-ray images can be produced by using two different monoenergetic or quasi-monoenergetic x-ray spectra. Two different monoenergetic x-ray beams at different energies will produce different x-ray images, because the two differentbeams will be transmitted and absorbed differently by the tissues and other substances in the patients body. By subtracting the two images, a differential x-ray image can be produced which will often emphasize or enhance the visibility of particular tissues or substances in the patients body. Thus, for example, the visibility of x-ray contrast materials can be enhanced. Examples of such materials are iodine, xenon and barium, which can be introduced into the bloodstream, the

lungs or the food canal of the patient.

Instead of using two different monoenergetic x-ray beams, itis generally more convenient to employ two different quasi-monoenergetic x-ray spectra, which may have peaks at different energy levels. The two different x-ray spectra will produce different x-ray images image. By using this technique, the visibility of contrast media or particular tissues can often be enhanced.

The two different quasi-monoenergetic x-ray spectra can be produced by using two different x-ray filters alternately. The x-rays to be filtered may be derived from an ordinary x-ray tube which produces a continuous spectrum of x-rays over a wide band of energies. The filters produce selective absorption of the x-rays so as to modify the continuous spectrum to produce peaks at different energy levels.

One important purpose of producing the differential x-ray images is to subtract or cancel out the portionsof the x-ray images which are due to the ordinary tissues of the patient, particularly the soft tissues, so that the visibility of certain distinctive tissues or contrast substances in the patients body will be enhanced. To produce good cancellation, the portions of the two different images which are due to the soft tissues must be balanced. However, even the ordinary soft tissues of the patient transmit x-rays of different energies to different extents. While the two different x-ray images can be balanced for any particular thickness of the patient, variations in such thickness over the field of view tend to upset such balance. Thus, without compensation for patient thickness variations, cancellation of the x-ray images due to the ordinary tissues of the patient can be achieved at only one value of patient thickness.

One object of the present invention is to provide effective compensation for variations in the thickness of the patient over the field of view. so that cancellation of the image portions dueto ordinary soft tissue can be accomplished to a high degree over a range of patient thicknesses.

A further object is to provide effective compensation for patient thickness variations without resorting'to a complex electronic compensation system.

In accordance with the present invention, compensation for patient thickness variations is achieved by adjusting the nature of the monoenergetic or quasi-monoenergetic x-ray spectra. It has been found that the quasi-monoenergetic spectra can be adjusted by adjusting the x-ray filtration and the supply voltage of the x-ray source. The adjustment of the voltage changes the x-ray spectrum produced by the x-ray tube. It is prefered to adjust the x-ray filtration so that one of the x-ray filters transmits a spectrum having two x-ray peaks at energies below and above the energy of the peak produced by the other x-ray filter. Thus, the image produced by the first x-ray filter consitutes the sum of the image components produced by the two peaks. Such sum tends to remain in a constant relationship to the image produced by the second spectrum, despite variations in patient thickness. The peaks of the x-ray spectra can be adjusted by varying the composition and density of the x-ray filters. It is prefered to vary the supply voltage to the x-ray tube as an inverse function of the average patient thickness. Thus, the voltage is reduced when the average patient thickness in creases. The adjustment of the supply voltage changes the x-ray spectrum produced by the x-ray tube, and thereby modifies the two different quasi-monoenergetic x-ray spectra transmitted by the two x-ray filters. By adjusting the supply voltage, it is possible to optimizethe compensation for patient thickness variations.

which can be subtracted to produce a differential x-ray Manydifferent materials may be employed in the x-ray filters. For example, one filter may contain iodine, while the other contains cerium. This combination of filters is particularly valuable for enhancing the visibility of iodine in the patients body. The iodine filter produces a spectrum having first and second peaks at low and high energy levels. The cerium filter transmits a spectrum having a peak at an intermediate energy, higher than the low energy peak but lower than the high energy peak produced by the iodine filter. By adjusting the supply voltage, the relative magnitudes of the high and low energy peaks can be adjusted.

Further objects, advantages and features of the present invention will appear from the following description, taken with the accompanying drawings, in which:

FIG. 1 is a diagrammatic illustration of an x-ray system to be described as an illustrative embodiment of the present invention.

FIG. 2 is a set of graphs illustrating the variation in the x-ray attenuation coefficients, as a function of x-ray energy, for iodine and for water, which accounts for most of the attenuation produced by soft tissue.

- FIG. 3 is a set of graphs illustrating the variation in the x-ray attenuation coefficients for iodine and for cerium'.

FIG. 4 is a set of graphs illustrating the wide band x-ray spectrum produced by an x-ray tube, and also the quasi-monoenergetic x-ray spectra which may be produced by using iodine and cerium filters in conjunction with the x-ray tube.

FIG. 5 is a set of graphs illustrating the effects of variations in the patient tissue thickness, with and without compensation.

FIG. 6 is a set of graphs illustrating the manner in which compensation can be achieved for different 4 ranges of patient thickness.

FIG. 7 is a graph illustrating the manner in which the x-ray supply voltage may be varied to achieve compensation for different values of patient thickness.

FIGS. 8-12 are additional graphs illustrating the effects of changing the filtration densities and the supply voltage.

As just indicated, FIG. 1 illustrates a system 10 for producing differential x-ray images, with compensation for variations in the thicknessof the patient or subject 12 to be x-rayed. In general, the thickness of the portion of the patient to be x-rayed may vary in a more or less irregular manner over the field of view. While the cross sectional shape of the patient 12 is shown as a simple oval, it will be understood that the actual shape is more or less irregular.

The x-ray system 10 comprises an x-ray source assembly 14 for producing monoenergetic or quasi-monoenergetic x-ray spectra. In this case, the source assembly 14 is adapted to produce quasi-monoenergetic x-ray spectra. Thus, the x-ray source assembly 14 comprises an ordinary x-ray tube or source 16 which may-be energized by a variable high voltage supply 18. An x-ray source of this type produces a continuous x-ray spectrum over a wide band of energies,

represented by a graph 20 in FIG. 4. The maximum energy of the band of x-rays is determined by themaximum voltage applied to the x-ray tube 16, in this case sixty-five kilovolts peak (KVp).

The x-ray beam from the x-ray source 16 is directed through the patient 12 to an image detector 22. To produce quasi-monoenergetic x-ray spectra, it is prefered to provide a plurality of selectively usable x-ray filters, two such filters 24a "and 24b being shown. The x-ray filters 24a and b may contain various materials capable of absorbing x-rays in a selective manner, so that the continuous band of x-rays produced by the x-ray tube 16 will be converted into x-ray spectra having one or more peaks at various energy levels. A wide variety of filtering materials may be employed. For example, the filters 24a and b may utilize iodine and cerium. FIG. 4 includes graphs 26a and b representing quasi-monoenergetic spectra which may' be produced by the use of the iodine and cerium filters. It will be seen that the spectrum 26a produced by the iodine filter 24a comprises a low energy peak 26aL and a high energy peak 26aH. The magnitude of the low energy peak 26aL is greater than the magnitude of the high energy peak 26aH. Similarly, the spectrum 26b produced by the cerium filter includes a low energy peak 26bL and a high energy peak 26bH. In this case, the high energy peak 26bI-I is so small as to be insignificant. The magnitude of the low energy peak 26bL for the cerium filter is comparable to .the magnitude of the low energy peak 26aL for the iodine filter. It will be seen that the low energy peak 26aL for the iodine filter occurs at a lower energy than the low energy peak 26bL for the cerium filter. On the other hand, the high energy peak 26aH for the iodine filter occurs at a higher energy than that of the low energy peak 26bL for the cerium filter.

The x-ray filters 24a and b are arranged to be used alternately, so as to produce two different x-ray images. As shown in FIG. 1, the filters 24a and b are movable alternately into the beam of x-rays from the 'x-ray tube 16. Thus, the filters 24a and b may be mounted on a movable indexing member, such as the illustrated rotatable disc 28. A mechanical device, such as the illustrated motor 30, may be employed to rotate the disc 28.

The two different x-ray images produced by the use of the filters 24a and b are detected by the image detector 22 and are subtracted by an image subtraction device 32. The differential images thus produced may be displayed in some suitable fashion, in this case by a television monitor 34.

Any known or suitable system may be employed for subtracting the x-ray images. For example, the image detector 10 may be arranged to produce positive and negative optical images corresponding to the two different x-ray images. The positive and negative images are then combined to produce the desired subtraction.

The prefered system is to convert the two different x-ray images into television images which are subtracted electronically to produce a differential television image, to be displayed on the television monitor 34. .Thus, the image detector 22 preferably comprises a television system for converting the x-ray images into electronic television signals. Theimage subtraction device 32 comprises means for subtracting the successive television images. Such means may utilize one or more image storage tubes, or other storage devices. For example, a silicon target storage tube may be employed, so that one of the two images can be written negatively, while the other image is written positively. In this way, any identical portions of the two images will be cancelled out. A differential image will develop on the silicon screen of the storage tube. This differential image can be displayed on the television monitor 34.

While various image subtraction systems may be employed, it is prefered to utilize the highly advantageous image subtraction system disclosed in the co-pending patent application of Charles A. Mistretta and Michael G. Ort, Ser. No. 369,824, filed June 14, 1973. Such system uses two stages of subtraction. In such system, the x-ray images are converted into video signals, which are then processed through two successive stages of video subtraction, preferably utilizing two different types of video. storage devices. The second stage of ages in the form of electrical charges distributed overa dielectric layer on a conductive target back plate.

The first and second video images are supplied sequentially to the storage tube, which produces a first differential video signal corresponding to the difference between the first and second video signals. The first differential video signal is then supplied to a second integrating and subtracting storage device, prefe rably utilizing a second storage tube capable of storing video images in the form of electrical charges on a mosaic of dielectric islands on a conductive back plate. The first differential video signal is written positively on the target of the second storage tube. The first storage tube isthen employed to develop a second differential video signal corresponding to the difference between the second and first video image signals. The second differential video signal is written negatively on the target of the second storage tube. In this way, the differential features of the first and second differential video signals are integrated and enhanced, while the identical or non-differential features of the first and second differential video signals are combined subtractively so as to cancel them from the target of the second storage tube.

To obtain the maximum enhancement of the differential features, this cycle of subtraction and integration is repeated so that a multiplicity of cycles are completed. The enhanced image on the target of the second storage tube can be read, as desired, and reproduced as a visible display on a television monitor. By this system, differential features amounting to only a fraction of one percent of the full contrast range of the successive x-ray images can be enhanced to full contrast so that such differential features will be clearly visible.

Reference may be had to such co-pending application for a more detailed disclosure of such differential imaging system. Any other known or suitable differential imaging system may be employed.

The iodine and cerium filters 24a and b are especially valuable for visualizing small quantities of iodine in the patients body. Iodine ispresent naturally, particularly in the thyroid gland. Moreover, iodine may be introduced into the patients body through the bloodstream or the food canal, to serve as a contrast medium or substance.

FIG. 2 comprises a graph 36 representing the variation of the x-ray attenuation coefficient for iodine, as a function of the x-ray energy, expressed in kiloelectron volts (KEV). It will be seen that the x-ray attenuation coefficient generally decreases with increasing x-ray energy. However, the graph 36 has an abrupt discontinuity, usually refered to as the k-edge 38, at which the x-ray attenuation coefficient increases sharply. Thus, at an x-ray energy E, above the k-edge energy, the attenuation coefficient is abruptly and substantially greater than at an x-ray energy E below the k-edge. The iodine and cerium filters make it possible to utilize the k-edge discontinuity 38 to produce an .enhanced differential feature between the two x-ray images, so that the visibility of such feature will be increased in the differential x-ray image.

Thus, it will be seen from FIG. 4 that the low energy peak 26a L for the iodine filter is at approximately 30,000 electron volts, below the k-edge energy of about 33,000 electron volts. The low energy peak 26bL for the cerium filter is at an energy of about 40,000 electron volts, above the k-edge energy. Thus, the x-ray spectrum produced with the cerium filter will be attenuated much more by iodine than will the spectrum produced with the iodine filter.

FIG. 3 comprises graphs which compare the x-ray absorption coefficients for iodine and cerium. It will be seen that FIG. 3 again includes the graph 36 for iodine, showing the k-edge discontinuity 38. FIG. 3 also includes a graph 40 for cerium. Here again, the x-ray absorption coefficient generally decreases with the increasing x-ray energy. However, there is a sharp discontinuity or k-edge 42 at a particular energy level, at which the x-ray absorption coefficient increases sharply. The k-edge 42 for cerium is at a substantially higher energy than the k-edge 38 for iodine.

The spectrum 26a for the iodine filter in FIG. 4 includes a sharp drop 44 which corresponds to the k-edge 38 for iodine. Similarly, the spectrum 26b for the cerium filter in FIG. 4 includes a sharp drop 46 which corresponds to the k-edge 42 for cerium.

FIG. 2 also includes a graph 47 which represents the variation of the x-ray attenuation coefficient for water, as a function of x-ray energy. It will be seen that attenuation coefficient decreases gradually with increasing x-ray energy. The absorption coefficient for water corresponds generally to the absorption coefficient for soft tissue. It will be seen that the absorption coefficient for the energy E, above the k-edge for iodine is slightly less than the absorption coefficient for the energy E below the k-edge 38. It is possible to compensate for this slight difference for any particular value of patient thickness. This can be done, for example, by adjusting the relative densities of the iodine and cerium filters 24a and b so that the intensity of the x-rays transmitted through the iodine filter 24a and the pertinent portion of the patients body is the same as the intensity of the x-rays transmitted through the cerium filter at the pertinent portion of the patients body.

However, in accordance with the present invention, it has been found that effective compensation can be provided for variations in the thickness of the patient over a wide range, by adjusting the filters and the supply voltage to the x-ray tube, so as to produce particularly advantageous quasi-monoenergetic x-ray spectra.

As previously indicated, one of the filters, in this case the iodine filter 24a, is adjusted to produce an x-ray spectrum having two quasi-monoenergetic peaks, while the other filter, in this case the cerium filter 26b, is adjusted to produce only one significant quasi-monoenergetic peak. The two peaks produced by the first filter are-below and above the energy level of the peak produced by the second filter. Thus, as to x-rays in the high energy peak for the first filter, the absorption coefficient for softbody tissues is less than for x-rays in the peak produced by the second filter.

As previously indicated, spectra of this type are illustrated in FIG. 4, in which the spectrum 26a for the iodine filter has a low energy peak 26a L and high energy peak or bump 26a H. At an intermediate energy, the spectrum 26b produced by the cerium filter has a single significant peak 26bL. There may be a small high energy peak or bump 26bH for the cerium filter, but this small bump is so insignificant as to be negligible.

Quasi-monoenergetic spectra of this 'type produce compensation for variations in the thickness of the patient, because the high energy peak 26aH. of the first filter gradually predominates over the intermediate energy peak 26bL of the second filter as the patient thickness increases, while the low energy peak 26aL of the first filter gradually predominates over the peak 26bL for the second filter, with decreasing patient thickness. Due to this action, it is possible to achieve a close balance between the x-ray intensities transmitted through the two filters and the patients body for a considerable range of patient thicknesses.

It will be evident from FIG. 4 that as the patient thickness increases, the low energy iodine peak 26aL is reduced with respect to the main cerium peak 26bL, but the high energy peak or bump 26aH of the iodine spectrum grows relative to the cerium peak 26bI-I. The sum of the high and low energy iodine portions 26aL and 26aI-I of the spectrum 260 thus remains approximately equal to the transmitted intensity of the cerium spectrum. The ratio of the transmitted intensities can be tuned to unity over a fairly broad range of patient thicknesses, as illustrated in FIG. 5, which compares the ratios of the transmitted intensities for the case of monoenergetic x-ray lines situated at the positions of the main iodine and cerium bumps 26aL and 26bL and for the case of three monoenergetic x-ray lines approximating the spectra provided by the iodine and cerium filters.

Thus, FIG. comprises a first graph 50, shown in a broken line, which plots the ratio of I to I as a function of tissue thickness, expressed in centimeters. IABOVE is the intensity of the x-rays transmitted through the tissue at an x-ray energy of 36,000 electron volts (keV). This energy is above the k-edge for iodine. I LOW is the transmitted x-ray intensity through the body tissue for an x-ray energy of 30 keV, below the k-edge for iodine.

It will be seen that the graph 50 rises steadily with increasing tissue thickness, and passes through unity for only one value of tissue thickness, approximately centimeters.

FIG. 5 includes a second graph 52, shown in a full line, in which IBELOW is produced by two x-ray spectral lines, at 30 keV and at 50 keV. The latter line, at 50 keV, corresponds to the high energy peak 26aI-I of the spectrum 26a produced by the iodine filter 24a. It will be evident that the graph 52 has a broad plateau 52a, centered at an average value of patient thickness which can be chosen by adjusting the filters and the supply voltage. In this case, the plateau 52a is centered at an average patient thickness of about 15 centimeters. This plateau provides a broad region of relative insensitivity to patient thickness. Thus, the patient thickness can vary over the field of view without materially affecting the ability of the differential x-ray system to cancel out the portions of the x-ray images due to ordinary soft tissue.

The exact location of the plateau 52a of FIG. 5 can be changed by'varying the filters and the voltage supplied to the x-ray tube. Suprisingly, it has been discovered that the center of the plateau moves to greater patient thicknesses as the supply voltage to the'x-ray tube is lowered. Thus, the supplyvoltage needs to be adjusted as an inverse function of the average patient thickness. The supply voltage needs to be decreased as the average patient thickness is increased, and vice versa. Thus, for example, it has been found that for average patient thicknesses of centimeters or more, it is highly desirable to use a supply voltage of 50,000 volts peak (kVp), with high tube current to produce the desired intensity of the x-rays, rather than raising the tube voltage.

Thus the supply voltage to the x-ray tube is an important factor in achieving effective compensation for patient thickness variations. The effect of changing the supply voltage is illustrated in FIG. 6, which comprises a series of graphs representing the ratio of cerium to iodine transmissions, as a function of patient thickness, for different supply voltages. The density of the cerium filtration was 200 milligrams per square centimeter.

Specifically, FIG. 6 comprises three graphs 54a, 54b and 54c, plotted for supply voltages of 50 kVp, 60 kVp and 70kVp. Each graph showsa plateau, which is inverted, in this case, because the ratio being plotted is the inverse of the ratio plotted in FIG. 5. The plateau of the curve 540, plotted for 50 kVp, is centered at a patient thickness of about 22 centimeters, while the curves 54b and 540, for supply voltages of 60 and kVp, are centered at patient thickness values of about 1 l centimeters and 6 centimeters, respectively. The transmission ratio at the center of each plateau is not unity, but compensation for this factor can be made by adjusting the electronic gain in the differential x-ray imaging system. Thus, different values of electronic gain can be used in the television system for the images produced with the use of the iodine and cerium filters, to bring about the optimum cancellation of the portions of the images representing ordinary soft tissue. The plateaus of the curves shown in FIG. 6 then provide wide regions in which the effectiveness of the cancellation is insensitive to variations in patient thickness.

FIG. 7 comprises a graph 56 in which the desirable x-ray supply voltage is plotted against patient thickness. This graph highlights the desirability of reducing the supply voltage to the x-ray tube as the average patient thickness is increased. By thus reducing the supply voltage, it is possible to optimize the compensation for variations in the patient thickness.

If the average patient thickness is decreased, the supply voltage should be increased to optimize the compensation for patient thickness variations. Thus, the supply voltage should beadjusted as an inverse function of the average patient thickness.

It is possible to change the composition and density of the x-ray filters so as to change the quasi-monoenergetic spectra produced by the use of the filters. For a particular average patient thickness, it is possible to adjust the filters so as to optimize the compensation for variations in the patient thickness.

As indicated by the legends in FIG. 4, the curve 26a represents a quasi-monoenergetic spectrum produced by the use of an iodine filter having a density or concentration of 0.200 grams of iodine per square centimeter. The quasi-monoenergetic spectrum represented by the curve 2612 was produced by a cerium filter having a density or concentration of 0.238 grams of cerium per square centimeter.

It is readily possible to vary the concentration of the iodine, cerium or other material in the filter. For example, this can be done by constructing each filter so as to include a tank or other receptacle which can be filled with a liquid solution containing the iodine, cerium or other material in a dissolved state. The concentration of the solution can readily be varied. The tank or receptacle may be made of a plastic material or some other material which produces very little attenuation of x-rays.

The curves of FIG. 6 were produced with cerium and iodine filters. The cerium filter has a density or concentration of 200 milligrams of cerium per square centimeter. The iodine concentration of iodine filter was cornparable and was adjusted to produce a ratio of unity between the iodine and cerium transmissions for zero patient thickness. The patient thickness was then varied for several different values of supply voltage to the x-ray tube, so as to produce the three curves 54a, 54b and 540.

In FIG. 6, a percentage figure is indicated by a legend for each curve, representing the percentage of the differential x-ray signal'for an iodine concentration of l milligram per square centimeter in the patient. The percentage signal applies to the broad minimum or inverted peak in the curve. Each percentage figure represents the maximum signal that can be obtained for that particular supply voltage to the x-ray tube. For a supply voltage of 50 kVp, the maximum differential signal was approximately 1 percent. For a supply voltage of 60 kVp, the maximum differential signal was approximately 0.95 percent. For a supply voltage of 70 kVp, the maximum differential signal was approximately 0.8 percent. It will thus be evident that the percentage value of the maximum differential signal decreases with increasing supply voltage to the x-ray tube. The graphs of FIG. 5 were obtained for monoenergetic spectra which approximate the quasi-monoenergetic spectra produced by iodine and cerium filters, as represented by Flg. 4. The curves 50 and 52 represent the results of calculations based on the use of monoenergetic xrays. Thus, the curve 50 represents the ratio of IABOVE to IBELOW as a function of tissue thickness, when the x-ray energy EABOVE is 36 keV, while the x-ray energy EBELOW is 30 keV. The x-ray energy E of 36 keV is above the k-edge for iodine, while the x-ray energy EABOVE of 30 keV is below the k-edge for iodine. The x-ray intensities [ABOVE and [BELOW represent the transmitted x-ray intensities for the x-ray energies EABOVE and EBELOW- The curve 52 of Flg. 5 represents the ratio of IABOVE to IBELOW for the same x-ray spectra, except that the x-ray component EBELOW at 30 keV is supplemented by an x-ray component at 50 keV having an intensity of approximately 14 percent of the intensity of the 30 keV component.

The x-ray component at 36 keV corresponds generally to the cerium peak 26bL of FIG. 4, which is above the k-edge of iodine, while the x-ray component of 30 keV corresponds generally to the iodine peak 26aL, which is below the k-edge for iodine. The x-ray. component at 50 keV corresponds generally'to the high en ergy bump 26al-I for iodine, as represented in FIG. 4.

FIGS. 8, 9 and 10 represent the results which are obtained by using different values of filter density or thickness, for different values of the supply voltage to the x-ray tube. In each graph, the ratio of TB to TA is plotted as a function of variations in the tissue thickness of the patient. TB is the transmission of x-rays through the patient with the use of the iodine filter, which has its major spectral peak below the k-edge for iodine. Such peak is designated 26aL in FIG. 4. TA is the transmission of the x-rays through the patient with the use of the cerium filter, having its 'major peak 26bL above the k-edge of iodine.

FIG. 8 comprises four curves 68a, 68b, 68c and 68d, representing the results obtained by using four different sets of filters at an x-ray supply voltage of 50 KVp. For the curve 68a, the cerium filter had a density of 0.100 grams of cerium per square centimeter, while the iodine filter had a density of 0.101 grams of iodine per square centimeter. For an iodine concentration in the patient of one milligram per square centimeter, the differential signal has a percentage value of 0.76 percent at a patient thickness of centimeters.

For the curve 68b, the corresponding values are as follows: cerium filter density 0.200; iodine filter density 0.195; percentage of differential signal, 1.19 percent.

For the curve 680, the corresponding values are as follows: cerium filter density 0.300; iodine filter density 0.284; percentage of differential signal, 1.42 percent.

For the curve 68d, the corresponding values are as follows: cerium filter density 0.400; iodine filter density 0.370; percentage of differential signal, 1.54 percent.

It will be understood that the symbol TA in FIGS. 8, 9 and 10 represents the total number of transmitted photons in the cerium spectrum, having its major peak above the k-edge for iodine. The symbol TB represents the total number of transmitted photons in the iodine spectrum, having its major peak below the K-edge for iodine. For each curve, the relative densities of the cerium and iodine filters were adjusted to produce a ratio of unity between TB and TA, for a patient tissue thickness of zero.

In each of the curves 68a-d of FIG. 8, the ratio of TB to TA decreases with increasing tissue thickness toward a minimum which occurs at approximately centimeters or greater, depending upon the density of filtration. These curves represent the situation for an x-ray tube supply voltage of 50 KVp. For this relatively low supply voltage, and for even lower voltages, any increase in the density of filtration tends to increase the tissue thickness at which the minimum occurs.

FIG. 9 comprises four curves 69a, 69b, 69c and 69d, which represent the ratio of TB to TA as a function of patient tissue thickness, for increasing values of filtration, at a supply voltage of 60 KVp. For the curve 69a, the cerium filter had a density of 0.100 grams per square centimeter, while the iodine filter had a density of 0.101 grams per square centimeter. The differential signal was 0.60 percent at a tissue thickness of 10 centimeters for an iodine concentration in the patient of one milligram per square centimeter.

The corresponding values for the curve 69b are as follows: cerium filtration 0.200; iodine filtration 0.197;

differential signal 0.95 percent.

For the curve 690, the corresponding values are as follows: cerium filtration 0.300; iodine filtration 0.288; differential signal 1.18 percent.

For the curve 69d, the corresponding values are as follows: cerium filtration 0.400; iodine filtration 0.376; differential signal 1.34 percent.

It will be observed that for the supply voltage of 60 KVp, represented by FIG. 9, the ratio of T8 to TA decreases with increasing tissue thickness until a minimum value is reached, whereupon the ratio increases with further increases in the tissue thickness. The increases in the filtration values do not increase the tissue thickness at which the minimum occurs, to any great extent. Thus, the position of the minimum remains at about 10.5 centimeters, despite the changes in the concentrations of cerium and iodine in the filters. The curves of FIG. 9 indicate that a supply voltage of 60 KVp is appropriate for an average tissue thickness of about 10.5 centimeters, because this value of supply voltage provides effective compensation for variations in patient tissue thickness. Such compensation is optimized by a supply voltage in this general neighborhood.

FIG. 10 comprises four curves 70a, 70b, 70c and 7011, representing the ratio of TB to TA as a function of pa tient tissue thickness for a supply voltage of 70 KVp, at four different filter concentrations or densities. For the curve 70a, the cerium concentration was 0.100 grams per square centimeter. The iodine filter concentration was 0.102 grams per square centimeter. The differential signal was 0.46 percent for an iodineconcentration of one milligram per square centimeter in the patient.

The corresponding values for the curve 70b are as follows: cerium concentration 0.200; iodine concentration 0.201; differential signal 0.70 percent.

The corresponding values for the curve 700 are as follows: cerium concentration 0.200; iodine concentration 0.298; differential signal 0.86 percent.

For the curve 70d, the corresponding values are as follows: cerium concentration 0.400; iodine concentration 0.392; differential signal 0.99 percent.

It will be observed that the curve 70a has a minimum at a tissue thickness of about seven centimeters. The ratio of TB to TA increases for smaller or larger values of tissue thickness.

With increasing filtration, the minimum value of the ratio tends to occur at smaller values of tissue thickness. Thus, the minimum value of the curve 70b is at a tissue thickness of about 6 centimeters. Thus, in general, at a supply voltage of 70 KVp or higher, increasing the filtration tends to decrease the tissue thickness at which the minimum occurs. Moreover, the value of the ratio of TB to TA at the minimum point tends to increase with increasing filtration. This is the opposite of the situation for supply voltages of 50 or 60 KVp, at which the value of the ratio at the minimum tends to decreasewith increasing filtration.

In summary, the effect of changing the densities or concentrations of the filters depends upon the supply filters does not affect the position of the minimum point to any substantial extent. At 50 KVp or lower, increasing the filtration tends to increase the tissue thickness at which the minimum ratio occurs.

For an average tissue thickness of about 10.5 centimeters, 60 KVp is an appropriate supply voltage, because the compensation for variations in the patient tissue thickness will be optimized. For average values of patient tissue thickness ranging down to 5 centimeters or less, the voltage should be increased progressively to 70 KVp or higher, to optimize such compensation. As the average tissue thickness increases to 20 centimeters or higher, the voltage should preferably be reduced progressively to 50 KVp or lower, so as to optimize the compensation for variations in the tissue thickness.

FIG. 11 and 12 will be helpful in explaining the effects represented by FIGS. 8l0. It will be seen that FIGS. 1 l and 12 reproduce the graph 36 of FIG. 3, representing the x-ray absorption coefficient for iodine, plotted as a function of x-ray energy. The x-ray absorption coefficient is designated m in FIGS. 11 and 12. As before, the graph 36 includes the abrupt k-edge 38.

FIGS. 11 and 12 also include the graph 26a of FIG. 4, representing the quasi-monoenergetic spectrum produced by the use of the iodine filter 24a. The graph 26a is superimposed upon the graph 36 in each case.

In FIG. 11, the spectrum graph 26a is drawn for an x-ray supply voltage of 50 KVp, but in FIG. 12, the graph 26a is drawn for a supply voltage of 70 KVp. In FIG. 11, m represents the value of the x-ray attenuation coefficient for iodine at the x-ray energy value corresponding to the low energy peak 26aL of the iodine filter spectrum 26a. On the other hand, m represents the value of the x-ray attenuation coefficient at the energy level corresponding to the high energy peak or bump 26aI-l of the iodine filter spectrum. It will be seen that m is substantially greater than m As a result, at a supply voltage of 50 KVp, the high energy bump 26aI-I loses to the low energy bump 26aL as the filters are made thicker or denser. Asa result of this decrease in the relative strength of the high energy bump or peak 26aI-I, a greater tissue thickness is required before the transmitted x-rays due to the high energy bump 26aH catch up with the x-rays transmitted by the cerium filter. The energy level of the x-rays transmitted by the cerium filter is substantially lower than the energy level represented by the high energy bump 26al-I, so that a greater percentage of the x-rays represented by the high energy bump are able to penetrate the increased tissue thickness.

The effect is just the reverse at KVp, as will be evidentfrom FIG. 12. It will be seen that the x-ray attentuation coefficient m at the energy of the low energy bump or peak 26aL is greater than the x-ray attenuation coefficient m at the x-ray energy corresponding to the high energy bump 26aI-I. Thus, as the filters are made thicker ormore dense, the xrays due to the high energy bump 26aH actually gain in magnitude relative to the x-rays due to the low energy bump 26aL. Thus, less tissue thickness is required for the more pentrating x-rays due to the high energy bump 26aH to equal or exceed the x-rays transmitted through the patient due to the cerium spectrum 26b, shown in FIG. 4. Thus, the minimum along each curve in FIG. 10 occurs at a relatively small tissue thickness. As the filters are made more dense, the minimum occurs at a decreased tissue thickness. This represents the situation at 70 KVp.

InFIG. 8, representing the situation at 50 KVp, the minimum along each curve occurs at a relatively great tissue thickness. As the filters are made more dense, the minimum occurs at a greater tissue thickness.

FIG. 9 represents an intermediate situation at 60 KVp. In this case, the minimum along each curve occurs at an intermediate tissue thickness. Changes in the tissue thickness do not affect the location of the minimum to any great extent.

From the graphs of FIGS. 8, 9 and 10, it will be possible for those skilled in the art to select an appropriate supply voltage and appropriate filter densities, according to theaverage tissue thickness involved, in order to achieve effective compensation for variations in the tissue thickness. The supply voltage and the amount of filtration should be selected to produce a curve having a minimum at or near the average tissue thickness.

It will be evident that iodine and cerium filters may be employed very advantageously for producing differential x-ray images due to iodine in the patients body. Such iodine may be present naturally, as in the thyroid, or may be introduced into the patients body as a contrast agent. The iodine and cerium filters are also valuable for producing differential x-ray images due to xenon gas, which may be inhaled into the lungs of the patient. Ordinary xenon gas can be employed, because the xenon does not have to be radioactive. The xenon ga's produces a k-edge which is at a somewhat higher energy level than the k-edge for iodine, but at a lower energy level than the k-edge for cerium. Thus, the iodine and cerium filters can be employed to produce a differential x-ray image due to xenon gas in the lungs of the patient. By this technique, the lungs can be visualized with a greater degree of clarity.

Various other contrast media may be employed to produce differential x-ray images. Barium is an example of another contrast substance. For each contrast substance, the composition and density of the x-ray filters 24a and 2417 are selected so as to produce quasi-monoenergetic x-ray spectra having peaks above and below the k-edge for the particular contrast substance. By proper selection of the x-ray supply voltage and the filter densities, effective compensation can be achieved for variations in the tissue thickness of the patient.

We claim: l. A method of producing differential x-ray images to visualize a contrast material in a patient,

comprising the steps of producing a first x-ray image using an 'x-ray source and first filter,

producing a second x-ray image using said x-ray source and a second filter,

said filters producing substantially different x-ray spectra,

producing a differential image corresponding to the difference between said first and second x-ray images,

said contrast material being visualized in said differential image,

and adjusting the voltage of said x-ray source as in inverse function of the average patient thickness to minimize the effect of variations in patient thickness upon said differential image.

2. A method according to claim 1,

in which the voltage of said x-ray source is adjusted to a range which produces a flat peakof the ratio between said first and second x-ray images when plotted against patient thickness.

3. A method of producing differential x-ray images,

comprising the steps of producing a first x-ray image using a first x-ray spectrum,

producing a second x-ray image using a second x-ray spectrum,

and producing a differential image corresponding to the difference between said first and second x-ray images,

said second spectrum having a spectral element of a predetermined energy,

said first x-ray spectrum comprising a low energy spectral element and a high energy spectral element,

said low energy spectral element having an energy below said predetermined energy,

said high energy spectral element having an energy above said predetermined energy.

4. A method according to claim 3,

in which said first and second x-ray spectra are balanced to provide substantially the same image elements in said first and second x-ray images for soft tissue whereby said image elements will cancel out in said differential image.

5. A method according to claim 4,

in which said spectral elements are adjusted to produce substantial balance between said first and second spectra over a substantial range of tissue thickness.

6. A method according to claim 3,

in which said first and second spectra are produced by using first andsecond x-ray filters in conjunction with an x ray source.

7. A method according to claim 6,

in which said first and second filters contain iodine and cerium, respectively.

8. A method according to claim 6,

in which said x-ray source includes a high voltage x-ray tube,

the supply voltage to said x-ray tube being varied as an inverse function of the average tissue thinkness to maintain a substantial balance between said first and second spectra for soft tissue over a substantial range of tissue thickness.

9. A method according to claim 6,

including the additional step of changing the density of said filters as a direct function of the average tissue thicknessof the patient.

10. A method according to claim 6,

including the steps of adjusting the density of said filters as a direct function of the average tissue thickness of the patient while adjusting the x-ray energy .of said source as an inverse function of said average tissue thickness to achieve effective compensation for variations in said tissue thickness.

11. A method according to claim 6,

including the steps of adjusting the density of said filters while adjusting the x-ray energy of said source to achieve effective compensation for variations in the tissue thickness of the patient.

12. A method according to claim 6,

including the steps of adjusting the density of said filters while adjusting the x-ray energy of said x-ray source to produce a flat spot in the characteristic curve of TB/TA plotted as a function of patient tissue thickness at a desired average value of patient tissue thickness,

where TB is the total of the x-ray photons transmitted through the first filter and the patient while TA is the total of the x-ray photons transmitted through the second filter and the patient,

whereby effective compensation is achieved for variations in the patient tissue thickness.

13. A method according to claim 12,

in which said first and second filters contain iodine and cerium, respectively.

14. A method according to claim 13,

in which said filters have iodine and cerium concen' trations on the order of to 400 milligrams per square centimeter.

15. A method according to claim 12,

inwhich said fiat spot is in the form of a minimum along said characteristic curve.

16. A method according to claim 12,

in which said x-ray source includes an x-ray tube supplied with a high voltage,

said voltage being adjusted as an inverse function of the patient tissue thickness.

17. A method according to claim 16,

in which said first and second filters contain iodine and cerium, respectively.

18. A method according to claim 17,

in which the iodine and cerium concentrations of the filters are on the order of 100 to 400 milligrams per square centimeter.

in which said high voltage is on the order of 50 to 70 kilovolts peak.

20. A method according to claim 17,

in which said high voltage is on the order of 50 to 70 kilovolts peak.

21. Apparatus for producing differential x-ray images in which image elements due to the soft tissue of a patient are at least partially cancelled out,

comprising x-ray source means for selectively producing first and second x-ray spectra,

means for producing first and second x-ray images of the patient using said first and second x-ray spectra,

and means for producing a differential image corresponding to the difference between said first and second x-ray images,

said second x-ray spectrum having an x-ray spectral element of predetermined energy,

said first x-ray spectrum having low and high energy spectral elements,

said low energy element having an energy less than said predetermined energy,

said high energy element having an energy greater than said predetermined energy.

22.- Apparatus according to claim 21,

in which said spectral elements have magnitudes such as to produce substantial balance between said first and second x-ray images for soft patient tissue.

23. A method according to claim 22,

in which said spectral elements are of magnitudes to maintain the balance between said first and second x-ray images for soft tissue over a wide range of tisin which said x-ray source means comprises an x-ray source and first and second selectively usable filter means for producing said first and second x-ray spectra.

25. Apparatus according to claim 24,

in which said first and second filter means are movable selectively into operative relation with said x-ray source.

26. Apparatus according to claim 24,

in which said first and second filter means contain iodine and cerium, respectively.

27. Apparatus according to claim 24,

in which said first and second filter means contain iodine and cerium, respectively, in concentrations on the order of 100 to 400 milligrams per square ccntimeter.

28. Apparatus according to claim 27,

in which said x-ray source includes a high voltage x-ray tube and means for supplying said tube with a high voltage on the order of 50 to kilovolts peak.

29. An apparatus according to claim 24,

in which said first and second filter means contain iodine and ceriui'n, respectively,

said x-ray source including a high voltage x-ray tube and means for supplying said tube with a high voltage on the order of 50 to 70 kilovolts peak.

30. Apparatus according to claim 24,

in which said x-ray source comprises a high voltage x-ray tube,

and means for varying the high voltage supplied to said tube.

Patent Citations
Cited PatentFiling datePublication dateApplicantTitle
US3588502 *Mar 3, 1970Jun 28, 1971George B GreenfieldRadiographic image subtraction technique
Referenced by
Citing PatentFiling datePublication dateApplicantTitle
US3965358 *Dec 6, 1974Jun 22, 1976Albert MacovskiCross-sectional imaging system using a polychromatic x-ray source
US3974386 *Mar 27, 1975Aug 10, 1976Wisconsin Alumni Research FoundationDifferential X-ray method and apparatus
US4110581 *Mar 12, 1976Aug 29, 1978C.G.R. MevDevice for monitoring the fitting in position of detachable elements in their support
US4204225 *Nov 15, 1978May 20, 1980Wisconsin Alumni Research FoundationReal-time digital X-ray subtraction imaging
US4204226 *May 16, 1978May 20, 1980Wisconsin Alumni Research FoundationReal-time digital X-ray time interval difference imaging
US4393402 *Jun 8, 1981Jul 12, 1983General Electric CompanySubtraction fluoroscopy method and apparatus
US4393512 *Jul 7, 1981Jul 12, 1983The United States Of America As Represented By The United States Department Of EnergyHyper-filter-fluorescer spectrometer for x-rays above 120 keV
US4409615 *Dec 29, 1980Oct 11, 1983Thomson-Csf Broadcast, Inc.Video imaging with automatically controlled radiation source
US4450478 *Sep 9, 1981May 22, 1984Georgetown UniversityDigital fluorographic method and system
US4499591 *Nov 17, 1982Feb 12, 1985Gary HartwellFluoroscopic filtering
US4511799 *Dec 10, 1982Apr 16, 1985American Science And Engineering, Inc.Dual energy imaging
US4683377 *Jul 5, 1983Jul 28, 1987Fuji Photo Film Co., Ltd.Subtraction processing method and apparatus for radiation images
US4806776 *Aug 5, 1985Feb 21, 1989Kley Victor BElectrical illumination and detecting apparatus
US4887604 *May 16, 1988Dec 19, 1989Science Research Laboratory, Inc.Apparatus for performing dual energy medical imaging
US4933960 *Jan 23, 1989Jun 12, 1990Yukio FujisakiFilter selection method for hard x-ray analysis of objects
US4947414 *Jul 14, 1986Aug 7, 1990Hologic, Inc.Bone densitometer
US5003571 *Jun 1, 1990Mar 26, 1991Nippon Identograph Co., Ltd.X-ray image equipment
US5033073 *Oct 24, 1988Jul 16, 1991Boeing CompanySystem for radiograhically inspecting a relatively stationary object and related method
US5044002 *Mar 7, 1989Aug 27, 1991Hologic, Inc.Baggage inspection and the like
US5148455 *May 28, 1991Sep 15, 1992Hologic, Inc.Bone densitometer
US5157703 *Jun 18, 1988Oct 20, 1992Wolf Hans DetlefMethod and device for the examination of cellular changes in an organism
US6075877 *Oct 21, 1997Jun 13, 2000Fuji Photo Film Co., Ltd.Method and apparatus of image processing using energy substraction parameters according to thickness of an object
US6226352Sep 8, 1998May 1, 2001Veritas Pharmaceuticals, Inc.System and method for radiographic imaging of tissue
US6330302 *Dec 22, 1999Dec 11, 2001U.S. Philips CorporationMethod of and device for forming an image of an object from plurality of images
US6370223 *Apr 6, 2001Apr 9, 2002Ut-Battelle, LlcAutomatic detection of bone fragments in poultry using multi-energy x-rays
US6714621 *Jul 6, 2001Mar 30, 2004Ge Medical Systems Global Technology Company LlcMethod and apparatus for radiological examination by injection of a contrast medium
US6723746 *Mar 15, 2001Apr 20, 2004Veritas Pharmaceuticals, Inc.Functional radiographic imaging methods and agents
US6735273Sep 27, 2001May 11, 2004Siemens AktiengesellschaftX-ray computed tomography apparatus and multi-spectra correction using a radiation pre-filter
US6751290Mar 16, 2001Jun 15, 2004Veritas Pharmaceuticals, Inc.Radiographic assessment of tissue after exposure to a compound
US6923950Dec 29, 2000Aug 2, 2005Veritas Pharmaceuticals, Inc.hexokinase ligands for diagnosis of cancer
US7397904May 11, 2005Jul 8, 2008Varian Medical Systems Technologies, Inc.Asymmetric flattening filter for x-ray device
US7649981 *Oct 15, 2003Jan 19, 2010Varian Medical Systems, Inc.Multi-energy x-ray source
US7869862Oct 15, 2003Jan 11, 2011Varian Medical Systems, Inc.Systems and methods for functional imaging using contrast-enhanced multiple-energy computed tomography
US7924970 *Jan 28, 2010Apr 12, 2011Siemens AktiengesellschaftMethod and device for generating a CT image with a high time resolution
US7949101 *Jun 16, 2009May 24, 2011Rapiscan Systems, Inc.X-ray scanners and X-ray sources therefor
US8311185Jan 7, 2010Nov 13, 2012Varian Medical Systems, Inc.Multi-energy x-ray source
US8503616Sep 24, 2008Aug 6, 2013Varian Medical Systems, Inc.X-ray tube window
US8625735Apr 14, 2011Jan 7, 2014Rapiscan Systems, Inc.X-ray scanners and X-ray sources therefor
US20110243382 *Feb 22, 2011Oct 6, 2011Edward James MortonX-Ray Inspection System and Method
USRE37536Mar 4, 1997Feb 5, 2002Uab Research FoundationSplit energy level radiation detection
EP0038663A1 *Apr 14, 1981Oct 28, 1981Technicare CorporationMethod of selecting a preferred difference image
EP0059382A1 *Feb 19, 1982Sep 8, 1982General Electric CompanySpinning filter for X-ray apparatus
EP0066805A2 *May 28, 1982Dec 15, 1982General Electric CompanySubtraction fluoroscopy method and apparatus
EP0162321A1 *Apr 24, 1985Nov 27, 1985Fuji Photo Film Co., Ltd.High-speed image recording apparatus for energy subtraction processing
EP0231037A1 *Jan 7, 1987Aug 5, 1987Philips Electronics N.V.X-ray scanner with dual energy imaging
EP2002287A2 *Mar 15, 2007Dec 17, 2008Philips Intellectual Property & Standards GmbHDynamic optimization of the signal-to-noise ratio of dual-energy attenuation data for reconstructing images
WO1988010094A1 *Jun 18, 1988Dec 29, 1988Hammann GeorgProcess and device for investigating cellular changes in a living organism
WO2000013590A2 *Aug 31, 1999Mar 16, 2000Jesse SalbSystem and method for radiographic imaging of tissue
WO2002026133A1 *Sep 27, 2001Apr 4, 2002Siemens AgX-ray computer tomograph
WO2012097801A1 *Jan 17, 2012Jul 26, 2012Bayer Pharma AktiengesellschaftMethod for generating a contrast medium-assisted x-ray image and x-ray system
Classifications
U.S. Classification378/62, 378/98.11, 378/108, 378/157
International ClassificationH05G1/60, A61B6/00, A61B6/03
Cooperative ClassificationA61B6/481, A61B6/504, A61B6/50, H05G1/60, A61B6/482, A61B6/4035, A61B6/00
European ClassificationA61B6/50H, A61B6/48D, A61B6/40F, A61B6/50, H05G1/60, A61B6/00