BONE MINERAL CAMERA

Abstract: An apparatus for determination of the mineral content in bones and comprising an X-ray source (12) capable of releasing X-ray radiation at at least two different energy levels for radiation of a preselected target area, and a camera part including a scintillator (22) receiving the X-ray radiation which have passed through the target area and emitting visible light, a CCD-sensor (25), and optical means for focusing the radiation from the scintillator on the CCD-sensor (25). The CCD-sensor communicates with an electronic circuit adapted to read out the CCD-sensor and to transfer data signals to a computer, said data signals representing the received radiation intensity and consequently a number of images (usually two received at their respective energy level) of the part of a person which should be examined. The apparatus is adapted to limit the energy content of the X-ray radiation emitted through the person in such a manner that the skin of said person is not subjected to more than a predetermined, allowed amount of energy. In this mananer an apparatus for determination of the mineral content in bones of living persons is provided, and this apparatus is smaller and less expensive than the hitherto known apparatuses at the same time as it is easy and quick to use, whereby the examination can be performed by a specialized physician.

Description: Title: Bone mineral camera- Technical Field The present invention relates to an apparatus for determination of the mineral content in bones and comprising an X-ray source capable of releasing X-ray radiation at at least two different energy levels for radi¬ ation of a preselected target area, a scintillator receiving the X-ray radi¬ ation having passed through the target area and emitting visible light, a CCD sensor, and optical means for focusing the radiation from the scintillator on the CCD sensor. A knowledge of the mineral content in bones is of vital importance for the possibility of evaluating a person's risk of being subjected to frac¬ tures and consequently for the possibility of evaluating the optional preventive treatment applying. A demand applies for a possibility for the specialized physicians of performing such a determination of the mineral content in bones because the possibilities of finding the patients needing the preventive treatment are thereby considerably increased. The energy content of the applied radiation should be limited such that the patient does not receive more than the maximum allowed amount of energy on his skin. Data signals representing a number of images (usually two) of the body area of the patient to be examined are trans¬ ferred from the CCD-sensor or the camera to a computer. Background Art WO 91 /09495 discloses an X-ray apparatus of the above type which images the entire measuring area on a CCD through a scintillator. US-PS No. 4,947,414 and the corresponding Danish Patent Application No. 1320/88 discloses furthermore an apparatus for determination of the bone density, said apparatus scanning the target. Similar appar¬ atuses have been described in several other publications, such as EP 0 253 742; EP 0 274 522; EP 0 432 730; and WO 91 /1 1 147. Such apparatuses are used together with a computer of a capacity sufficiently high for allowing processing of image signals. The computer includes programmes allowing it to both process data signals represent¬ ing the X-ray images and based thereon to produce images of the bone density in a target. Furthermore, the computer is preferably adapted to adjust the calculations in response to periodic calibration measurements on a known target in such a manner that the intensity or the colour of each image element, the so-called pixels, contains information on the density of the tissue through which the radiation has passed. The com¬ puter can also be adjusted to calculate and display the density either in each pixel or for a small group of pixels, such as four, or for a larger area, such as one vertebra. The apparatuses used today are all very cost-intensive and bulky and are only found in hospitals. The known apparatuses used today are scanners employing a narrow X-ray beam or a fan beam moved across the target area of the patient and combined with a correspondingly moved detector. The object of the invention is to provide an apparatus for determination of the mineral content in bones in living persons, where the apparatus employs less radiation doses and is less expensive than the hitherto known apparatuses at the same time as it is easy and quick to use such that the examination can be performed by a local physician. Description of the Invention The apparatus according to the invention is characterised in that the X- ray source is arranged to emit a short pulse of high energy radiation, the X-ray source being adapted to be applied a voltage of about 90 kV during insertion of an Au-filter in the path of radiation, whereby the emitted radiation downstream the Au-filter is of an average quantum energy of about 75 ke V, and that the X-ray source is further arranged to emit a low-energy radiation (38 keV) of a slightly longer duration than the first pulse shortly after or before said first-mentioned radiation, the X-ray source being applied a voltage of about 50 kV during insertion of a Ce-filter in the path of radiation, whereby the emitted radiation down¬ stream the Ce-filter is of an average quantum energy of about 38 keV. The two elements Au and Ce are particularly suited for the determination of the bone density because their atomic configurations provide them with the capacity of absorbing and consequently filtering off energies exceeding 88 keV (Au) and 40.5 keV (Ce), respectively, with a very sharp cut off. Through a suitable selection of density and X-ray acceler¬ ation voltage it is thereby possible to form narrow X-ray energy spectres of about 4 keV 3 dB-width. The necessary density is about 5 mg/mm 2 for both elements. The X-ray acceleration voltage depends slightly on the X-ray tube in question, but is in the range of 10 to 15% above the desired average energy. As a result it is possible to produce absorbtion images at energies which are favourable for an accurate determination of the content of calcium. This provides in turn improved measuring results as well as a possibility of using a low dose of radiation. Since any X-ray radiation must be considered injurious to health on a long view, it is of vital importance that the radiation dose is low, and especially in connection with diagnostic apparatuses used for prophylac¬ tic screening examinations. In practise the gold filter is a thin layer of gold sheet in form of a disk of a diameter of about 15 mm placed close to the X-ray source (the source being substantially punctiform). The most economic way of providing the cerium filter is to admix ce- riumoxide (which is a conventional optical abrasive) in epoxy in suitable amounts and to cast a thin epoxy disk of a diameter of for instance 15 mm. The apparatus operates as a camera, i.e. the image is taken in one step whereas the data are read out in series to the computer quite like in the known and hitherto used scanner. The apparatus is less complicated than the scanners as all the parts are permanently mounted and immov¬ able during the exposure. Although CCD-sensors are slightly sensitive to both X-ray radiation and visible light, a considerably improved detection of the X-ray radiation is obtained by the camera part comprising a scintillator in addition to the CCD-sensor, said scintillator receiving the X-ray radiation through the examined area of the person for each exposure. After each exposure with X-ray radiation, the scintillator emits electromagnetic radiation of a longer wavelength in form of light towards the CCD-sensor. The scintillator may be a plastic film with crystals embedded therein of a light-emitting substance, such as CaWO 4 , Y 2 O 2 S:Tb and/or Gd 2 O 2 S, and preferably about 300 mg/cm 2 of Gd 2 O 2 S or more being highly sensi¬ tive to said types of radiation. In principle, the scintillator and the CCD-sensor could be arranged abut- ting one another in such a manner that the emission of light from the scintillator would occur exactly in front of the CCD-sensor. It is, how¬ ever, desired to operate with a reasonably large target area on the patient of about 18 cm times 12 cm, which implies that the detection area should preferably be about 24 cm times 16 cm. It is difficult and cost-intensive to produce CCD-sensors in this size by means of the present technology, and although it is possible to provide a sufficiently large detection area by employing several rows of CCD-sensors, a draw¬ back is found in the read-out period for data from all these CCD-sensors being unexpediently long. Accordingly, the camera part comprises an optical lens means imaging the scintillator on the light-sensitive surface of the CCD-sensor. The imaging ensures a reduced image in such a manner that the entire image can be displayed on a CCD-sensor of a suitable size, such as 26 mm times 17.3 mm, whereby the data-reading can be performed very quick¬ ly, such as in a few seconds. According to a preferred embodiment, the CCD-sensor is cooled in order to reduce the dark current. Moreover, the camera part may advantageously comprise a mirror ab¬ sorbing X-ray radiation and reflecting light, and be mounted in such a manner that the light from the scintillator is emitted towards the optical lens means optionally through one or more additional mirrors, whereby a compact structure of the optics of the apparatus is obtained. Description of the Drawings The invention is explained in greater detail below with reference to some Examples and the accompanying drawings, in which Fig. 1 illustrates a highly simplified diagramme of an Example of an apparatus according to the invention, Fig. 2 illustrates an outline of the radiation path from the scintillator to the CCD-sensor through a preferred embodiment of a lens system for an apparatus according to the invention, Fig. 3 illustrates an Example of a structure of a lens system, Fig. 4 illustrates the distribution function of the probability of electrons in an electron well (pixel) on a CCD-sensor, Fig. 5 illustrates an Example of an alternative embodiment of an appar¬ atus according to the invention, Fig. 6 corresponds to Fig. 2, but with a concave mirror instead of one of the plane mirrors, and Fig. 7 illustrates an Example of an embodiment of an apparatus accord- ing to the invention, Fig. 8 is a front view of a further particularly preferred embodiment of an apparatus in use with the X-ray camera in a vertical camera position, Fig. 9 is a side view of the same embodiment, Fig. 10 is a top view of the same embodiment, Fig. 1 1 A and 1 1 B illustrate yet another outline of the apparatus in use with the X-ray camera in a horizontal camera position, Fig. 12 illustrates a change from vertical to horizontal camera position. Detailed Description The bone mineral scanner comprises a radiation source 12, which may be a small X-ray tube of the type used by dentists. The radiation source may for instance be arranged about 75 cm above the patient, and the detection equipment is arranged about 25 cm below, cf. Fig. 7. It is, of course, within the scope of the invention to arranged the radiation source below the patient and the detection equipment above the patient as indicated in Fig. 1. A couch must be provided for the specific examin¬ ation, such as the gynaecological couch of a specialized physician. The patient must be able to lie conveniently relaxed, preferably and usually substantially plane on his back, optionally merely recumbent. The equipment must, of course, be mounted such that the X-ray radi- ation path extends substantially perpendicular to the bones to be exam¬ ined. The X-ray tube 12 must operate at various voltages and with various filters. The following voltages are preferably used: 1 ) The radiation is emitted through an Au-filter, a voltage of 90 kV being applied to the tube, whereby the emitted radiation presents an average quantum energy of 75 keV downstream the Au-filter. 2) The radiation is emitted through a Ce-filter, a voltage of 50 kV being applied to the tube, whereby the emitted radiation presents an average quantum energy of 38 keV downstream the filter. The radiation is emitted only in a very short period. For each of the above-mentioned voltages the exposure periods are in the range of about 0.1 to 1 sec depending on the thickness of the patient. The apparatus is preferably adapted to interrupt the radiation source when the radiation dose at the detector has reached a predetermined level so that the radia- tion dose is kept constant at the detector. Radiation sources of this nature are known technique and are not described in greater detail here. The patient must be placed on a couch, cf. for instance Fig. 7 showing a "Hammock-couch" or a gynaecological couch, so that the radiation is emitted through the examining target area of the patient. Typically, it is necessary to examine an area of about 20 cm times 15 cm. It is import- ant that the patient is placed in a convenient position so that he does not change his position during the exposure. The camera part is placed below the patient, said camera part preferably comprising a collimator 21 , a scintillator 22, a mirror 23, some optics 24, and a CCD-camera 26 or CCD-sensor 25, and an A/D-converter communicating with a computer 28. As illustrated in Fig. 1 , the data signals are transferred to a computer 28, said data signals being read out from the CCD-camera 26. Pro¬ grammes already exist and are available for processing such data so that an image is produced and the mineral content in the exposed bones is calculated, and accordingly they are not described in greater detail here. The X-ray radiation source 12, an associated voltage supply 14, and the camera part 22 to 26 are suitably provided on a movable frame, cf. for instance Fig. 7. Thus the entire equipment can be easily moved forwards into position above a suitable couch when an examination of this type is required. The distance from the X-ray anode to the patient is for instance 75 cm, and the distance from the patient to the detector can be 25 cm. An examination of an area of the patient of 180 mm times 120 mm requires a detection area of 240 mm times 160 mm. Fig. 2 illustrates an Example of the optical system of an apparatus accor- ding to the invention. The radiation is emitted from the scintillator 22 into a first mirror 23 provided on the back with a coating for instance of a laminate of Cu, Sn, and Pb absorbing the X-ray radiation which has continued through the scintillator 22. An ionizing chamber controlling the exposure period can be coupled between the mirror and the absorber. The mirror 23 reflects the green light converted by the scintillator towards an additional mirror 33, from which the light is collected by a lens system 34. A CCD-sensor 25 is mounted below the lens system, said sensor being cooled on its bottom side by Peltier- elements 30 in turn being cooled by a fan 31 . The preferred embodiment employs a CCD type 05-20 from EEV presenting very large pixels (22.5 μm times 22.5 μm) and suitable outer dimensions (26 mm times 17.3 cm). Fig. 3 illustrates a possible embodiment of a lens system made of 8 lenses providing an aperture ratio of 1 :0.714. Many options apply, how¬ ever, for combining various lenses, and the invention is not limited to the lens system shown. As the measuring of the bone mineral is an examination which is to be performed repeatedly on a large group of people, the allowable radiation dose per patient by each examination is very low, viz. of the magnitude 50 microsieverts as surface dose adjacent the X-ray tube. Therefore the efficiency of the detector used for detecting the radiation having passed through the patient should be so high that almost any X- ray quantum passing through the patient presents a significant contribu¬ tion to the complete measuring result. In connection with a lumbar examination of an average person, the radiation must pass about 15 g of soft tissue and about 1 g of bone mineral per cm 2 perpendicular to the radiation direction. When passing through soft tissue of a mass absorption coefficient of about μ/p = 0.28 cm 2 /g, the 38 keV-radiation is attenuated inside the patient to about 1 .5% of the original intensity, and when passing through soft tissue and lumbar vertebras of a mass absorption coeffi¬ cient of about μlp = 1 .1 cm 2 /g, typically to 0.5% of the original inten¬ sity being limited to the previously mentioned maximum allowed radia¬ tion dose of 50 μS (microsievert). When passing through soft tissue of a mass absorption coefficient for the hard radiation of μlp = about 0.18 cm 2 /g, the 75 keV-radiation is attenuated typically to 1 .5% of the original intensity, and when passing through lumbar vertebras of a mass absorption coefficient of μlp = about 0.3 cm 2 /g, to about 5% of the original intensity. Since the 38 keV-radiation is attenuated approximately 10 times as much on its way through the patient as the 75 keV-radiation, the best measuring result is obtained by allowing the 38 keV-radiation to contri¬ bute with 45 microsieverts and the 75 keV to contribute with 5 micro¬ sieverts to the maximum allowed surface radiation dose of a total of 50 microsieverts, i.e. the radiation received by the patient on his surface adjacent the X-ray tube. The number N of 38 keV X-ray quantums allowed to pass per cm 2 into the patient has been limited to: N = 45 E-9 J/g = 3 E7crτv 2 38 keV • 1.6 E-16 J/keV ■ 0.28 cm 2 /g Of these about 0.005 ■ 3 E7 = 150000 will hit the scintillator per cm 2 in shadow of lumbar vertebras, and about three times more on areas lying only behind soft tissue. In connection with thin patients, the radiation dose adjacent the X-ray tube can be halved every time the thickness of soft tissue is reduced by 2.5 cm. In connection with fat patients, the surface radiation dose must be correspondingly increased. The radiation having passed through the patient is detected by a scintillator. In front, the scintillator can comprise a so-called lead screen which is a converging collimator 21 focusing the X-ray anode, and it is provided for absorbing radiation Compton-spread in the patient. The lead screen is followed by the actual scintillator 22, and it can be a plastic film with crystals embedded therein of a light-emitting substance, such as CaWO 4 , Y 2 O 2 S:Tb and/or Gd 2 O 2 S. At present, a sheet with Gd 2 O 2 S- crystals is preferred, such as a "Lanex fast" from KODAK. The density of this scintillator of a type sold today is 150 mg/cm 2 , and it absorbs about 40% of each of the two radiation energies. A scintillator with 300 mg/cm 2 Gd 2 O 2 S is, however, preferably used. The Gd 2 O 2 S-crystals result in a sheet possessing both a high X-ray radiation absorption and a high light yield by the X-ray energies used. It is, however, within the scope of the invention to use other sheets converting X-ray radiation into light which is electronically detectable. The present scintillator type converts the individual X-ray quantums into a large number of green light quantums of the wavelength 545 nm and the quantum energy 2.27 e V, and it operates with a degree of efficiency of 19% which is particularly high for scintillators. In other words, an X- ray quantum with a quantum energy of 38 keV absorbed in the scintil¬ lator causes emission of 38000 eV/2.27eV 0.19 = 3200 green light quantums emitted in the following milliseconds. After the absorption of an X-ray quantum, the intensity of the green light decreases exponential- ly with a time constant of 3 ms. In order to obtain a suitable image of columna lumbaliis or colium femoris, the scintillator must be about 24 cm times 16 cm. The area of the light detectors commercially available at a reasonable price is, how¬ ever, small compared to the area of the scintillator, which for instance of the "Large Area CCD"-type 05-20 by EEV is no more than 26 mm times 17.3 mm distributed on 1 152 times 770 electron wells. It can therefore be advantageous to include a lens system imaging the light emission surface of the scintillator on the light-sensitive surface of the CCD. The outer diameter of the lens system should be dimensioned in accord¬ ance with the diameter of the light detector and the aperture ratio of the objective, and the aperture ratio should be chosen as large as possible, i.e. as large as allowed in view of the lens errors and the costs involved. An optics with a focusing distance f = 50 mm, which is the standard for objectives, is suitable, whereas the aperture ratio should at least be 1 : 1.2, and more preferred 1 :1 and most preferred even larger. The scintillator measures 240 mm times 160 mm, whereas the active area of the CCD measures 26.0 mm times 17.3 mm, and accordingly the image must be reduced in the ratio of the image diagonals d π /d 2 = 9.24. The distance a from the scintillator to the optics, and b from optics to CCD are found by 1/a -_- 1/b = 1 /f and a/b = d. ] /d 2 resulting in a = f • (1 + d.,/d 2 ) = 50 mm • (1 + 9.24) = 512 mm. An aperture ratio of 1 :1.2 results in a diameter of the diaphragm of: d b1 = 50 mm/1 .2 = 41 .7 mm, and the angle below which the scintillator detects the diameter of the optics is: v = 2 arctg (41 .7 mm / 2 1 512 mm) = 4.66°. The light from the scintillator is emitted according to a Lambert distribu¬ tion, and as far as small angles about the perpendicular are concerned the relative light yield is: q = 4(1 - cos . v/2) = 0.33%. Correspondingly, the following light yields in connection with large aperture ratios are obtained: Aperture ratio of 1 : 1 v = 5.59° and q = 0.48% Aperture ratio of 1 :0.9 v = 6.21 ° and q = 0.59% Aperture ratio of 1 :0.714 v = 7.82° and q = 0.93% The lens system must collect as much of the emitted green light as possible and image the scintillator plate on the light-sensitive surface of the CCD. The company "Mikroglass" sells an optics of an aperture of 1 :0.714 resulting in v = 7.82° and q = 0.93%, and a 38 keV X-ray quantum contributes with 8 electrons to the pixel. An Example of such a lens system appears from Fig. 3, where the syst¬ em is made of a total of 8 lenses 41 to 48. The radiation path is shown to 3 image dots placed in the plane back of the last lens 48. According to a preferred embodiment, the lowermost, last lens 48 is made of quarts, and the CCD-sensor 25 is placed as close as possible to this last lens 48. The space between the lens 48 and the CCD 25 is filled with an inert gas, such as xenon, presenting a low thermal con- ductivity. The CCD 25 must preferably be cooled by means of a Peltier- element 30, and it should be arranged horizontally below the lens 48 in order to avoid convection. The lens system can, however, be structured in many ways, and the shown Example is merely included to illustrate a possibility. An interest- ing variant is to place a holographic lens preferably close to the scintillator. The surface of the CCD should be chosen as large as possible, and on this surface an image is formed of the X-ray radiation falling on the scintillator. Then the CCD must be read out and its data transferred to a computer, such as a personal computer used anyway by the physician. Two methods different in principles can be used for reading out a CCD: A fast method only measuring the reset level at the beginning of each line and then the measure levels pixel by pixel in fast sequence through the remaining line, and a slow method integrating the voltage a suitable period in each pixel, about 2 to 10 microseconds, for both the reset and the mea¬ suring level, and then digitizing the difference between the two levels with an accuracy of 12 to 16 bis. Depending on the integration period, the reading out takes 5 to 30 microseconds per pixel, as it is possible to perform the analogue-digital- converting parallel to the level measurements. The fast reading-out method presents an acceptable differential, but a poor absolute measur¬ ing accuracy. By the fast method, a CCD can be read out in 1/50 second, which is required for a video display, but the method can only be used for images observed visually. It is not suited for collecting data which later on are to form part of calculations, and where a digital representation with a high accuracy is required. For this purpose, the slow differential measur¬ ing of pixel by pixel must be used. An EEV CCD type 05-20 having 1 152 times 770 pixels results in a reading out of all the pixels in 887040 times 6 microseconds = about 5 seconds. The size of the pixels measured on the 240 mm times 180 mm large scintillator is 0.21 mm times 0.21 mm which is suited for an X-ray image, but unnecessarily fine-meshed for a BMC-image (Bone Mineral Content). In the apparatus according to the invention it is possible before the integration and the analogue-digital-conversion on the pixels of the CCD to add up the charges in two times two pixels, whereby the read-out time is reduced from about 5 to about 1.5 second, and the pixel size on the scintillator is increased to 0.42 mm times 0.42 mm. Such a pixej ^ size is suitable and the read-out time acceptable for BMC measurements as the patient must not move at all between the exposures by means of the low and the high X-ray quantum energy. Light amplifiers In connection with some embodiments it can optionally be advantageous to insert an electrostatic light amplifier between the scintillator and the CCD. Due to the vacuum achievable in practise the hitherto known amplifiers result, however, in formation of at least one heavy positive gas-ion per millisecond, and due to the high voltage in the amplifier these ions cause strong dot-wise electron emissions from the photocathode. The result is heavily exposed pixels distributing statically, i.e. they are not the same during the measuring on the patient and on "flood field", and accordingly it is impossible to correct this error. Dynamics and noise The read-out sensitivity of a CCD is below 10 electrons per "electron well" or pixel, and the individual pixels can contain up to 150,000 elec¬ trons. This provides a dynamic area of a factor 15,000 = 2 14 . In order to completely utilize this dynamics, it is necessary to digitize with at least 14 bits. During both the exposure and the reading out, thermal electrons, the so-called dark current, are fed to the pixels. The average value of this current is compensated for by performing an additional exposure of the same duration, but without the X-ray radiation, and by subtracting the result from the actual measurements pixel by pixel. The measuring results are, however, encumbered with a statistic uncer¬ tainty corresponding to the square root of the total number of thermal electrons. The number of thermal electrons and consequently the statis¬ tic measuring uncertainty can be reduced by cooling the CCD. It is for instance known to cool astronomic CCDs which are to be exposed for long periods and then read out by long integration periods for each pixel, by means of liquid nitrogen to -196°C. A simple Peltier-cooling easily providing about -60°C suffizes for the object of the invention. Whenever the temperature is reduced by 8°C, the number of thermal electrons is halved . A lowering of the temperature of 80°C reduces therefore the number of thermal electrons to one thou- sandth and consequently the statistic measuring uncertainty by a factor 30. Since all the pixels do not possess quite the same efficiency, and since the optical system is not quite linear, it can be necessary to correct the images with respect to uniformity. This is easily done by pixel by pixel to divide the images by a uniformly exposed image. Computer pro¬ grammes already exist for dividing the images by a standard image in connection with computer processing bone mineral measurements. Beyond being sensitive to visible light, a CCD is also sensitive to X-ray radiation. The radiation detector is therefore preferably structured such that the radiation passing the scintillator does not hit the CCD. The above can for instance be performed by deflecting the light from the scintillator about 90° in a mirror coated on the back with a laminate of Cu, Sn, and Pb, said coated mirror simultaneously absorbing the non- utilized X-ray radiation. According to the preferred embodiment, the mirror is a plane mirror, but it can within the scope of the invention also be a concave mirror provid¬ ing a converging radiation path. Then the succeeding lenses must be correspondingly adjusted in such a manner that a correct formation of image of the scintillator surface is obtained on the surface of the CCD- sensor. Statistics The probability or the possibilities of obtaining a sufficiently accurate measurement of the bone mineral content of a patient are stated below, where the allowed radiation dose per patient and the efficiency of the components forming part of the system have been considered. The available "Lanex fasf'-scintillator sold by KODAK results in an im¬ mediate loss of 60% of the X-ray quantums having passed through the patient. Accordingly, a scintillator of a density twice as high or more as the above scintillator is used, which results in a loss of only 40% which is acceptable. A condition for the apparatus being applicable is that each individual 38 keV X-ray quantum being absorbed in the scintillator must contribute significantly to the total measuring result, i.e. result in a measurable number of electrons in the electron wells of the CCD. A 38 keV X-ray quantum absorbed in the scintillator emits 3,200 green light quantums. An optics with a possible aperture ratio of 1 :1 emits 0.48% or 15 of these light quantums into the CCD. Approximately 75% of the area of the CCD are light-sensitive with a quantum efficiency of 36%. Consequently, each X-ray quantum contributes with an average of 4 electrons to the electron wells. It should here be recalled that the company "Mikroglass" sells an optics with an aperture of 1 :0.714. Such an optics results in v = 7.82° and q = 0.93%, and a 38 keV X- ray quantum contributes with 8 electrons to the pixel. The above number is encumbered with a high statistic uncertainty. The uncertainty is illustrated in Fig. 4 showing as an Example the probability P n in order to obtain exactly n electrons in a pixel when the number has been Poisson-distributed with an average value of 8: P n = 8 n /n! • e "8 . In other words, each absorbed X-ray quantum provides a contribution, which, however, is of almost the same magnitude for the same quantum energy. The statistic uncertainty applies, however, also to the reading out of the number of electrons and to the number of X-ray quantums falling inside each pixel - here 2 times 2 pixels in the CCD. A scintillator with 300 mg of Gd 2 O 2 S/cm 2 results in an absorption of about 60% of both 38 keV and 75 keV X-ray quantums, said quantums being converted into 3200 and 6300, respectively, green light quan¬ tums. A pixel size of 0.42 mm times 0.42 mm has the effect that within each pixel shadowed by a lumbar vertebra of an average person the following is absorbed: about 150.000 cm "2 ■ 0.60 • 0.42 2 cm 2 = 150 X-ray quantums of each type. The 38 keV quantums will in the CCD result in 8- 150 = 600 electrons/pixel = 0.1 fC/pixel, the 75 keV quantums will in the CCD result in 16- 150 = 1200 electrons/pixel = 0.2 fC/pixel. Within areas only covered by soft tissue, the following is absorbed: about 450 pieces of 38 keV quantums and about 225 pieces of 75 keV quantums. Within the CCD they will each provide 3600 electrons/pixel = 0.6 fC/pixel. Now the most critical number of electrons are discussed in detail, i.e. the electrons deriving from the 150 pieces of 38 keV X-ray quantums. They are absorbed within each pixel shadowed by lumbar vertebras. These electrons have been Poisson-distributed and are accordingly en- cumbered with a spreading of TδO = about 12 X-ray quantums. In the CCD 12 pieces of 38 keV quantums provide 8 times 12 « 100 electrons, and this number is high compared to both the statistic reading out uncertainty of about 10 electrons and to the statistic uncertainty of V 8- V 150 = about 35 electrons representing an average number of 1200 electrons deriving exactly from 150 X-ray quantums. A considerably larger and extremely cost-intensive CCD (of the type used by astronomers) renders it possible to obtain an additional reduc¬ tion of the neasuring uncertainty of 1 - /I 00 2 / Λ 00 2 + 10 2 + 35 2 = 6% or is utilized to obtain a reduction of the radiation dose of: 1 - 100/(100 + 10 + 35) = 30%. A reduction of this magnitude is insignificant. Another way of reducing the radiation dose is to use a scintillator pre- senting a higher density than the above 300 mg Gd 2 O 2 S, an optics of a higher aperture ratio than 1 :0.7 and/or optionally a coat¬ ing of the CCD in order to turn it more sensitive to green light. It is furthermore within the scope of the invention to provide an optics which is more efficient with respective to light, such as for instance an optics emplying concave mirrors for collecting the light from the scintillator, cf. for instance Fig. 6. CCD-sensor Many possibilities apply to the choise of CCD, such s a photon-counting "intensitifed"-CCD or a modified "slow-scan"-CCD. The latter type of CCD includes so-called "diluted" CCDs of a quantum efficiency of more than 80% and a reading-out noise of a few electrons, as well as a very high charge-transfer efficiency in connection with small electric charges. By using for instance very thin CCD exposed from the back it is possible to additionally double the number of electrons provided by a 38 keV X- ray quantum to the pixel, i.e. from 8 to 16 electrons, as this type of CCD possesses a quantum efficiency being twice as high for green light as the conventional CCDs. The preferred embodiment employs a CCD type 05-20 by EEV exactly meeting the requirements due to its very large pixels (22.5 μm times 22.5 μm) and optimum outer dimensions (26 mm times 17.3 mm). As the charges from two times two pixels are added up in the presently preferred embodiment and as the measuring results are later on divided by the measuring results of a flood field, it is of no importance whether relatively many pixels are missing provided said missing pixels are uni¬ formly distributed and do not concentrate in lines or lumps. Accordingly without involving a significant reduction of the efficiency it is possible to use CCDs which have otherwise been discarded and which are therefore inexpensive. An apparatus of the above structure and as explained above can be used in the following manner: The apparatus is connected to a computer, such as a personal computer, already used by the physician in his prac- tice. The computer must comprise programmes for processing the data signals (image signals) transmitted from the CCD camera. The pro¬ grammes must include routines for calibration and correction of measur- ing data on the basis of control measurements performed at regular intervals on known measuring targets. In this manner an essential meas¬ uring accuracy and reproducability is ensured because the evaluation by the physician of the measurements is often based on a comparison of measurements performed at intervals of months or years. At the examination, the patient is for instance placed as shown in Fig. 7 when the examination is to be directed towards the spine. The patient is briefly exposed initially to a radiation of 75 keV and then to a radiation of 38 keV. After each exposure, data are read out from the CCD-sensor to the computer. The patient must not move during or between the two exposures. By means of already known programmes the computer can then display an image on the screen of the examined area of the patient, whereby the colour or the light intensity in each image pixel represents a measur- ement of the density on the spot in question. The known programmes can furthermore calculate the density of selected areas of the image, such as corresponding to the individual lumbar vertebras. The advantage of the apparatus according to the invention is that it is relatively simple compared to the known apparatuses for the same pur- pose because it comprises no movable parts and because the exposure can be performed in a few seconds, said exposure only being performed briefly at the two different energy levels. Consequently, the apparatus is so simple that it can be used in the practice of local specialized phys¬ icians. An Example of an advantageous embodiment of an apparatus according to the invention is shown in function in Fig. 8. The patient to be exam¬ ined is placed as shown in Fig. 8 on a couch which can be moved into a desired position, whereby the relevant area of the patient is placed in the best possible manner relative to the apparatus, i.e. aligned between the X-ray tube and the scintillator screen of the camera part. The apparatus is preferably built into a C-shaped casing or frame 16, cf. Figs. 9 and 10, in form of a hollow tube (preferably of steel or alumin- ium and made of standard tubes and tube bendings) mounted on a frame or, as shown, on a suspension plate 19 secured directly on a wall in such a manner that the frame can be displaced and turned between two permanent camera positions, viz. a vertical and a horizontal camera position, for instance by means of an optionally motor-driven rack and pinion. In the position shown by means of a solid line in Fig. 9, an X-ray tube 12 is built in an oil chamber at the top. The associated electronics 15, such as high-voltage circuits and control electronics, is preferably mounted inside the tube 16 forming the casing. A marking light source 17 is arranged opposite the X-ray source 12 in a circumference just outside the output opening of the X-ray tube. The marking light source marks the target area of the patient. The Au and the Cerium filters are not shown, but are preferably mounted close to the X-ray source in such a manner that they can be inserted in the radiation path. The camera part 20 is placed at the other end of the casing 16 vertically below the X-ray tube 12. The camera part comprises a scintillator screen 22 followed by a mirror shield 23 and optics 24 as well as CCD 25 and electronics 27 for collecting data for the image being produced on the CCD. The camera part 20 must be connected to an edb monitor 28. The distance between the X-ray tube 12 and the scintillator screen 22 is preferably about 1 m. The scintillator screen 22 should be placed as close as possible just below the couch 50 of the patient. Figs. 1 1 A and B show the apparatus in use with the casing 16 in the horizontal position, i.e. a position in which the radiation path of the X- ray radiation emitted through the patient is substantially horizontal. Figs. 1 1 A and B show furthermore that the C-shaped casing 16 or frame is preferably mounted on a suspension plate 19, which is pivotal in such a manner that the apparatus can be pivoted towards the wall when not used. Fig. 12 shows how the position is changed. The previous vertical posi¬ tion is here indicated by a dotted line, the arrow P indicates the moving direction of the casing 16. The X-ray tube 12 is placed farthest away from the wall and the control knobs 18 for the apparatus can advan¬ tageously be placed on the side of the X-ray tube part 12. A supporting wheel 29 can be mounted on the camera part 20, said wheel rolling on a design cut edge of the suspension plate 19, said edge of the plate being cut so as to support the camera part while the position is changed and while the camera is in the vertical position.

Claims: Claims. 1. An apparatus for determination of the mineral content in bones and comprising an X-ray source capable of releasing X-ray radiation at at least two different energy levels for radiation of a preselected target area, a scintillator receiving the X-ray radiation having passed through the target area and emitting visible light, a CCD sensor, and optical means for focusing the radiation from the scintillator on the CCD sensor, c h a r a ct e r i s e d in that the X-ray source is arranged to emit a short pulse of high energy radiation, the X-ray source being adapted to be applied a voltage of about 90 kV during insertion of an Au-filter in the path of radiation, whereby the emitted radiation downstream the Au- filter is of an average quantum energy of about 75 ke V, and that the X- ray source is further arranged to emit a low-energy radiation (38 keV) of a slightly longer duration than the first pulse shortly after or before said first radiation, the X-ray source being applied a voltage of about 50 kV during insertion of a Ce-filter in the path of radiation, whereby the emitted radiation downstream the Ce-filter is of an average quantum energy of about 38 keV. 2. An apparatus as claimed in claim 1, characte rised in that the scintillator (22) is a plastic film with crystals embedded therein of a light-emitting substance, such as CaWO , Y 2 O 2 S:Tb and/or Gd 2 O 2 S, and preferably comprises about 300 mg/cm 2 of Gd 2 O 2 S. 3. An apparatus as claimed in claim 1 or2, characterised in that the X-ray source and the camera part comprising the scintillator, optics, and CCD-sensor are built into a common substantially C-shaped casing (16) in such a manner that they form a unit. 4. An apparatus as claimed in claim 3, characterised in that the unit comprising the X-ray source and the camera part is suspended in such a manner that it is displaceable between a horizontal position in which the X-ray path extends substantially horizontally, and a vertical position in which said X-ray path extends substantially vertically. 5. An apparatus as claimed in one or more of the claims 1 to 4, c h a r a c t e r i s e d in that the CCD-sensor (25) is mounted on a cooling element, especially a cooling element in form of a Peltier element, preferably a Peltier element dimensioned to cool the CCD-sen¬ sor to about -50 to -70°C. 6. An apparatus as claimed in one or more of the claims 1 to 5, cha racterised in that the X-ray radiation-sensitive unit in addi¬ tion comprises at least one mirror (23) absorbing X-ray radiation and reflecting light and mounted such that the light from the scintillator (22) is emitted towards the optical lens means (24). 7. An apparatus as claimed in claim 6, characte rised in that one or more of the mirrors (33) is/are a concave mirror of a curvature causing a convergence of the radiation from the scintillator (22). 8. An apparatus as claimed in claim 6 or 7, cha racterised in that at least two mirrors (33, 34) are inserted between the scintillator (22) and the optical lens means (24). 9. An apparatus as claimed in one or more of the claims 1 to 8, cha ra cte ri sed in that the CCD-sensor is of the diluted type of a high quantum efficiency. 10. An apparatus as claimed in claim 3 or 4, cha racterised in that the electronic circuits of the X-ray source and the camera part are arranged in whole tube lengths and tube bends together forming a sub¬ stantially C-shaped frame (16).

Inventors: BAERENHOLDT OLAF, NIELSEN STIG PORS, BAERENHOLDT, OLAF, NIELSEN, STIG, PORS
Applicants: BAERENHOLDT OLAF, HOLOMED APS, NIELSEN STIG PORS
Application year: 1993
Main classification: A61B6
IPC codes: A61B6, H04N5, H05G1