Patent Document

CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is based upon and claims the benefit of priority from the prior Japanese Patent Application No. 2000-007815, filed Jan. 17, 2000, the entire contents of which are incorporated herein by reference. 
     BACKGROUND OF THE INVENTION 
     This invention relates to an X-ray computer tomography apparatus, and more particularly to a thin-type X-ray computer tomography apparatus (hereinafter, referred to as an X-ray CT apparatus) provided with a small-sized high-voltage transformer capable of supplying power and stepping up the voltage by itself in a noncontacting manner. 
     In the field of X-ray CT apparatuses, high-speed tomographic techniques for turning the rotatable gantry section with respect to the static gantry section at high speed are being developed rapidly. With the high-speed tomographic techniques, not only a larger amount of image information can be obtained in a short time but also the time during which the subject is tied down. Thus, the techniques are very effective in a group medical examination as well as an ordinary physical examination. In recent years, a high-speed rotation of about 0.5 second has already been put to practical use and further an ultrahigh-speed rotation of less than 0.3 second is gradually gaining practicality. 
     Because of the problem of centrifugal force acting on the rotatable gantry section by high-speed rotation, the gantry of the X-ray CT apparatus have been required to be made smaller in size, particularly thinner along the axis of the body of the subject. Although various measures have been taken to improve the method of arranging the units in the rotatable gantry section, it is necessary to make each unit as small as possible to reduce the size of the rotatable gantry section itself. In the X-ray CT apparatus, the top on which the subject has been laid is inserted into a cylindrical space formed within the rotatable gantry section and then pictures are taken. In such an X-ray CT apparatus, it is necessary to improve the accessibility of the subject as in the magnetic resonance imaging apparatus. An improvement in the accessibility enables various medical procedures when the subject is inserted into the apparatus. Moreover, the subject has a less feeling of confinement in the cylindrical space. 
     The power supply from the static gantry section to the rotatable gantry section is carried out in a contacting manner or a noncontacting manner. One example of supplying power in a contacting manner is achieved by using a slip ring mechanism. As is well known, the slip ring mechanism has a brush provided on the rotatable gantry section and causes the brush to come into contact with the slip ring provided on the static gantry section, thereby supplying power from the static gantry section to the rotatable gantry section. The slip ring mechanism is considered to be unsuitable for high-speed rotation, because the friction between the brush and the slip ring produces heat and abrasion powder. In addition, since there is a possibility of electric discharges, the slip ring is regarded as unsuitable for power transmission of such a high voltage, for example, 10 kV or more as is applied across both ends of an X-ray tube. Under these conditions, several concepts of X-ray CT apparatuses that supply power from the static gantry section to the rotatable gantry section in a noncontacting manner have been proposed. 
     One known noncontacting-type X-ray CT apparatus is disclosed in U.S. Pat. No. 4,912,735. This X-ray CT apparatus supplies power in a noncontacting manner by electromagnetic induction. 
     FIG. 11 is a schematic circuit diagram of a conventional X-ray CT apparatus that supplies power from the static gantry section to the rotatable gantry section. FIG. 12 shows the location of the individual component parts. In FIG. 12, an AC/DC converter  14   b  is connected to an alternating-current (a.c.) power source  11  provided on the side face of the lower part of the inside of the static gantry section  111 . The output terminal of the AC/DC converter  14   b  is connected to an inverter  15 . The output of the inverter  15  is connected to the primary coil  116  of the static gantry section  111 . The primary coil  116  is wound around the cylindrical static gantry section  111  in such a manner that it surrounds the outer surface of the static gantry section  111 . The rotatable gantry section  112  has a cylindrical shape as the static gantry section  111  does and is provided on the static gantry section  111  on the same central axis of the cylinder in such a manner that it can rotate. On the rotatable gantry section  112 , a secondary coil  119  is provided in a position facing the primary coil  116  of the static gantry section  111 . Like the primary coil  116 , the secondary coil  119  is wound around the rotatable gantry section  112  in such a manner that it surrounds the outer surface of the rotatable gantry section  112 . A high-voltage transformer  113  is connected to the secondary coil  119 . A rectifier  20  is connected to the output terminal of the high-voltage transformer  113 . An X-ray tube  21  is connected to the output terminal of the rectifier  20 . A magnetic field generated at the primary coil  116  induces power at the secondary  119 . The electromagnetic induction enables power to be supplied from the static gantry section  111  to the rotatable gantry section  112 . 
     The conventional X-ray CT apparatus with the above configuration that supplies power in a noncontacting manner has the following problem. 
     As compared with an ordinary transformer where the cores are integrally formed, the leakage inductance between the cores of the separate primary coil  116  and secondary coil  119  is greater, impeding a high-frequency operation, which makes it difficult to miniaturize the unit. The miniaturization is possible only when the operation of the unit is carried out at higher speed. For this reason, in general, to reduce the leakage inductance, the primary coil  116  is arranged as close to the secondary coil  119  as possible or the coil windings are wound even in the grooves of the cores, thereby improving the degree of coupling. In such a manner of improving the degree of coupling, there arises a problem that realizing the high frequency operation by overcoming the leakage inductance between the separated cores is limited. Moreover, it makes high-voltage insulation difficult from the viewpoint of manufacturing techniques. Thus, to obtain a high voltage of about 75 kV to 150 kV on the secondary side, it is necessary to provide an additional high-voltage transformer  113 . This puts significant restriction on the rotatable gantry section being made smaller and thinner. In the X-ray CT apparatus, the primary coil  116  is wound around the cylindrical static gantry section  111  in such a manner that it surrounds the outer surface of the static gantry section  111  and the secondary coil  119  is wound around the rotatable gantry section  112  in such a manner that it surrounds the outer surface of the rotatable gantry section  112 , with the result that the distance between the windings facing each other is relatively long. This makes the parasitic capacitance large, making a high-frequency operation difficult, which is one of the causes of the difficulty in making the unit smaller and thinner. 
     Another known noncontacting X-ray CT apparatus is disclosed in Jpn. Pat. Appln. KOKAI Publication No. 7-204192 and Jpn. Pat. Appln. KOKAI Publication No. 8-336521. Each of these conventional X-ray CT apparatuses has the following configuration. 
     Each of the X-ray CT apparatuses comprises electromagnetic induction transmission means including a first winding provided on the fixed frame of a scanner and a second winding provided on the rotary section of the scanner in such a manner that it faces the first winding, and a high-voltage generator connected to the electromagnetic induction transmission means. Each of the X-ray CT apparatuses supplies specific power in a noncontacting manner by electromagnetic induction. 
     Each of these conventional X-ray CT apparatuses is provided with an additional high-voltage transformer to obtain a high voltage, because of the problem of the leakage inductance, as explained in the X-ray CT apparatus disclosed in U.S. Pat. No. 4,912,735. This is a serious hindrance in making the rotatable gantry section smaller and thinner. 
     Each of U.S. Pat. No. 5,105,351 and U.S. Pat. No. 5,272,612, which are assigned to the same assignee as the present invention, discloses a device for applying a high voltage to an X-ray tube. Each of these two devices includes a plurality of high-voltage transformers, taking the size reduction of transformers into account. However, there is no reference to noncontacting power supply by electromagnetic induction or a concrete application of the inventions to an X-ray CT apparatus. 
     BRIEF SUMMARY OF THE INVENTION 
     The object of the present invention is to provide a thin-type and small-sized X-ray computer tomography apparatus. 
     According to the present invention, there is provided an X-ray computer tomography apparatus comprising: a static gantry section; a rotatable gantry section which is provided on the static gantry section in a rotatable manner and has an X-ray tube for generating X rays; a frequency converting circuit which is connected to an alternating-current power source and converts the output voltage from the alternating-current power source into a desired high-frequency voltage; a high-voltage transformer which transmits the output of the frequency converting circuit from the static gantry section to the rotatable gantry section and steps up the output to a desired high voltage; and a rectifier circuit which converts the alternating-current voltage outputted from the high-voltage transformer into a direct-current voltage and supplies the direct-current voltage to the X-ray tube, wherein the high-voltage transformer including: a primary-side which is provided on the static gantry section and to which the output of the frequency converting circuit is supplied, and a resonance circuit with a capacitor connected to a winding of a secondary-side, which is provided on the rotatable gantry section and generates the high voltage. 
     Additional objects and advantages of the invention will be set forth in the description which follows, and in part will be obvious from the description, or may be learned by practice of the invention. The objects and advantages of the invention may be realized and obtained by means of the instrumentalities and combinations particularly pointed out hereinafter. 
    
    
     BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING 
     The accompanying drawings, which are incorporated in and constitute a part of the specification, illustrate presently preferred embodiments of the invention, and together with the general description given above and the detailed description of the preferred embodiments given below, serve to explain the principles of the invention. 
     FIG. 1 shows the configuration of a circuit for supplying power from the static gantry section to the rotatable gantry section in an X-ray CT apparatus according to a first embodiment of the present invention; 
     FIG. 2 is a sectional view showing the location of the individual parts inside the static gantry section and rotatable gantry section; 
     FIG. 3 shows a first modification of the circuit configuration of FIG. 1; 
     FIG. 4 shows a second modification of the circuit configuration of FIG. 1; 
     FIG. 5 shows a third modification of the circuit configuration of FIG. 1; 
     FIG. 6 shows a fourth modification of the circuit configuration of FIG. 1; 
     FIG. 7 shows a fifth modification of the circuit configuration of FIG. 1; 
     FIG. 8 shows a detailed configuration of the transformer rotating section in the rotatable gantry section in an X-ray CT apparatus according to a second embodiment of the present invention; 
     FIG. 9 shows the location of the transformer fixing section on the static gantry section and the transformer rotating section on the rotatable gantry section which face each other; 
     FIG. 10 is a perspective view showing the arrangement of the primary side core and secondary side core in a separate-type high-voltage transformer; 
     FIG. 11 shows the configuration of a circuit for supplying power from the static gantry section to the rotatable gantry section in a conventional X-ray CT apparatus; and 
     FIG. 12 is a sectional view showing the location of the individual parts inside the static gantry section and rotatable gantry section in the conventional X-ray CT apparatus. 
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Hereinafter, referring to the accompanying drawings, embodiments of the present invention will be explained. 
     FIG. 1 shows the configuration of a circuit for supplying power from the static gantry section to the rotatable gantry section in an X-ray CT apparatus according to a first embodiment of the present invention. FIG. 2 is a sectional view showing the location of the individual parts inside the static gantry section and rotatable gantry section. 
     As shown in FIG. 11 the X-ray CT apparatus of the first embodiment is composed of a static gantry section  12  and a rotatable gantry section  22  roughly divided as shown by broken lines. The static gantry section  12  includes an alternating-current (a.c.) power generator section  13  composed of an a.c. power source  11 , an AC/DC converter  14 , and an inverter  15  and a transformer fixing section  51 . The rotatable gantry section  22  includes a transformer rotating section  52 , a rectifier  20 , and an X-ray tube  21 . 
     The transformer fixing section  51  provided on the static gantry section  12  and the transformer rotating section  52  provided on the rotatable gantry section  22  constitute a separate-type high-voltage transformer  50 . The transformer fixing section  51  includes a primary coil  16  and a primary side core  17 . The transformer rotating section  52  includes a secondary coil  19  and a secondary side core  18 . The primary side core  17  is not formed integrally with the secondary side core  18 . The arrangement of the primary side core  17  and secondary side core  18  will be explained in a second embodiment of the present invention. 
     The AC/DC converter  14  is connected to the output terminal of the alternating-current (a.c.) power source  11  serving as an input power source. A plurality of inverters  15  are connected in parallel with the output terminal of the AC/DC converter  14 . The output terminal of each of the inverters  15  is connected to the primary coil  16  of the transformer fixing section  16 . The AC/DC converter  14  converts the a.c. voltage from the a.c. power source  11  into a direct-current (d.c.) voltage. The d.c. voltage is then supplied to the inverter  15 , which converts the d.c. voltage into a high-frequency a.c. voltage. 
     The reason why a plurality of inverters  15  are used in FIG. 1 is to prevent the whole of the X-ray CT apparatus from stopping the operation if one of the inverters  15  fails. By selecting the troubled inverter  15  and stopping it, the power can be controlled roughly. For the convenience of design, only one inverter  15  may be provided. The high frequency a.c. power generator  13  may have another configuration, as long as it generates power of desired frequency, for example, about 100 kHz. 
     The output of the high frequency a.c. power generator  13  is connected to the primary coil  16 . When a plurality of inverters  15  are used, such as this embodiment, the output of each of the plurality of converters  15  is provided to the respective primary coils  16 . Alternatively, if only one inverter  15  is used, the output of the inverter  15  is parallelly connected to a plurality of primary coils  16 . 
     As shown in FIG. 2, the static gantry section  12  is mounted on a base  60 . Near the static gantry section  12 , the a.c. power source  11  is provided. The static gantry section  12  has an opening  101 . Along to the cylindrical direction on outside of the opening  101 , the inverter  15 , primary coil  16 , primary side core  17 , and others are arranged. A doughnut-like disk rotatable gantry section  22  with an opening  100  in it is provided on the static gantry section  12  in such a manner that it can rotate continuously. The top  120  is inserted into the opening  101  of the static gantry section  12  and the opening  100  of the rotatable gantry section  22 . 
     Outside the opening  100  of the rotatable gantry section  22 , the secondary side core  18 , secondary coil  19 , rectifier  20 , and others are arranged. The X-ray tube  21  and X-ray detector  32  are provided on the rotatable gantry section  22  in such a manner that they face each other with the opening  100  between them. 
     The primary coil  16  is wound around almost the central part of the primary side core  17 . Two primary coils  16  may be wound around one primary coil  17 . 
     The primary coil  16  and the primary side core  17  are arranged around the static gantry section  12  so that the magnetic flux generated at the primary side core  17  may be supplied to the rotatable gantry section  22 . 
     The shape of the primary  17  is not limited to the squared-U shape. As long as the magnetic flux generated at the primary side core  17 , together with the secondary side core  18  arranged so as to face the primary side core  17 , can form a magnetic circuit, the primary side core may take another shape. 
     The squared-U-shaped secondary side core  18  facing the primary side core  17  is arranged around the ringed rotatable gantry section  22  placed so as to surround the top  120  of the couch, as is the primary side core  17 . 
     In this case, too, the secondary side core  18  may take another shape, as long as it, together with the primary side core  17 , can form a magnetic circuit. 
     The secondary coil  19  is wound around almost the central portion of the secondary side core  18 . 
     As shown in FIG. 1, a capacitor  24  is connected in series with the secondary coil  19 . The secondary capacitor  24  is designed to resonate with the inductance of the secondary coil  19 . The inductance of the secondary coil  19  makes the impedance higher as the frequency increases, which is one of the factors that hinder the high-frequency operation most. When a suitable value of the secondary capacitor is selected, the secondary impedance can be adjusted by resonance, which enables a high-frequency operation. 
     In the conventional X-ray CT apparatus in each of FIG.  11  and FIG. 12, only one secondary coil  19  was used. Since the resonance voltage at the secondary coil  19  is overhigh than the necessary output voltage of 10 kV or higher, it is technically difficult to produce an insulting of the secondary coil  19 . Furthermore, it is technically difficult to produce a capacitor capable of withstanding such a high voltage by itself. To overcome this drawback, the first embodiment uses a plurality of secondary coils  19 , thereby lowering the voltage generated in each secondary coil  19 . This makes it technically easy to realize an insulting of the secondary coil  19  or the capacitor  24 . 
     The series circuit of the secondary coil  19  and capacitor  24  is connected to the rectifier  20 . 
     Although the number of rectifiers  20  is the same as that of capacitors  24 , either the number of rectifiers  20  or that of capacitors  24  may be larger than the other. The rectifiers  20  rectify high frequency a.c. power into d.c. power. The secondary side core  18 , secondary coil  19 , secondary capacitor  24 , and rectifier  20  on the rotatable gantry section  22  constitute a high-voltage unit  23 . 
     The one-side ends of the high-voltage unit  23  are connected in series and similarly its other-side ends are connected in series. The resulting one end and other end are connected to one end and the other end of the X-ray tube  21 , respectively. The high-voltage unit  23 , X-ray tube  21 , and X-ray detector  32  are provided around the rotatable gantry section  22 , taking weight balance into account. In the conventional example, when the rotatable gantry section was rotating, the weight of the rotatable gantry section was large and developed a great centrifugal force of, for example, about 13 G, which was a factor preventing a high-speed rotating operation. 
     In the first embodiment, however, the series resonance on the secondary side enables a high-frequency operation, for example, an operation at 100 kHz, which helps make the inverter  15 , primary coil  16 , primary side core  17 , secondary coil  19 , and secondary side core  18  smaller and lighter. The smaller, lighter secondary coil  19  and secondary side core  18  particularly decrease the weight and space of the rotatable gantry section  22  remarkably. Since the secondary coils  19  and the secondary side cores  18  are circularly and evenly arranged on the rotatable gantry section, the section excels at rotation balance. 
     Also, since the capacitor  24  is provided for resonance, and certain degree of the leakage inductance of the secondary coil  19  is used for the construction of the resonance circuit, there is no need to take into account the leakage inductance of the secondary coil  19  as inhibition factor. 
     Therefore, the secondary coil  19  can be wound around the secondary side core  18  with a sufficient insulting distance between them. As high a voltage as 150 kV can be generated, making it unnecessary to provide an additional high-voltage transformer for generating a high voltage on the rotatable gantry section  22 , which enables the rotatable gantry section  22  to be made smaller and thinner remarkably. 
     Because the decreased number of component parts on the rotatable gantry section makes room for the space of the rotatable gantry section, it is possible to realize an X-ray CT apparatus with multiple tubes. Use of an X-ray CT apparatus with multiple tubes can improve time resolution of acquired image. 
     Hereinafter, various modifications of the circuit for supplying power from the static gantry section to the rotatable gantry section in the X-ray CT apparatus according to the first embodiment will be explained. 
     In a first modification of the first embodiment in FIG. 3, the capacitor  24  is connected in parallel with the secondary coil  19 . The capacitor  24  resonates with the leakage inductance of the secondary coil  19 . 
     In a secondary modification of the first embodiment in FIG. 4, a capacitor  31  is inserted in series between the output of the high voltage a.c. power generator section  13  and the primary coil  16 . In the second modification, the capacitor  24  is provided so as to resonate with the leakage inductance of the secondary coil  19  and the primary capacitor  31  is provided so as to resonate with the inductance of the primary coil  16 . Even when resonance not only on the secondary side but also on the primary side make the operating frequency higher, the primary capacitor  31  and secondary capacitor  24  can be selected according to the resonance, which enables a high-frequency operation. That is, the primary-side leakage inductance can be used effectively in the secondary modification. 
     In a third modification of the first embodiment, the capacitor  31  is inserted in series between the output of the high frequency a.c. power generator  13  and the primary coil  16  and the capacitor  24  is connected in parallel with the secondary coil  19 . In a fourth modification of the first embodiment, the capacitor  31  is inserted in parallel between the output of the high frequency a.c. power generator  13  and the primary coil  16  and the capacitor  24  is connected in parallel with the secondary coil  19 . Furthermore, in a fifth modification of the first embodiment in FIG. 7, the capacitor  31  is inserted in parallel between the output of the high frequency a.c. power generator  13  and the primary coil  16 . 
     Hereinafter, a second embodiment of the present invention will be explained. 
     FIG. 8 shows a detailed configuration of the transformer rotating section on the rotatable gantry section in an X-ray CT apparatus according to the second embodiment. FIG. 9 shows the location of the transformer fixing section on the static gantry section and the transformer rotating section on the rotatable gantry section which face each other. FIG. 10 is a perspective view showing the arrangement of the primary side core and secondary side core in a separate-type high-voltage transformer. 
     A plurality of high-voltage unit blocks, for example, as shown in FIG. 8, four high-voltage unit blocks B 1  to B 4  are arranged on the gantry section  22  to form a circumference as a whole. These blocks are connected electrically to each other by connectors C 1  to C 4 . One high-voltage unit block includes, for example, four high-voltage units  23 , m 1  to m 4 . This divided structure facilitates the replacement of the high-voltage units  23 . Such a divided structure may be applied to the transformer fixing section of the static gantry section  12 . 
     As shown in FIG. 9, the primary coil  16  and primary side core  17  on the static gantry section  12  are provided so as to face the secondary coil  19  and secondary side core  18  on the rotatable gantry section  22 . The shape of and the number of the secondary side cores  18  are so determined that all of the plurality of primary side cores  17  never fail to the secondary side cores  18 , even when the rotary section (rotatable gantry section  22 ) rotates. 
     The spacing between the primary side core  17  and the secondary side core  18  is about 1 mm. Note that the spacing is not limited 1 mm. 
     As shown in FIG. 10, the primary coil  16  is wound on the primary side core  17  whose cross section perpendicular to the direction of rotation of the rotatable gantry section  22  is shaped like an almost squared U. The primary coil  16  is wound on the central part of the squared-U shape of the primary side core  17 . The primary side core  17  is so positioned that the two ends of the squared U may face the rotatable gantry section  22  and the straight line connecting the two ends be perpendicular to the direction of rotation of the rotatable gantry section  22 . The direction in which the rotatable gantry section  22  rotates at that time is shown by a thick arrow. 
     At the ends of the squared-U shape of the primary side core  17 , there are provided two projecting sections  71  of the same shape which project in the direction opposite to the direction of rotation of the rotatable gantry section  22 . The projecting sections  71  may be made of the same material as that of the primary side core  17  or of a magnetic substance made of a material with different susceptibility. One projecting section  71  may be spaced, for example, about 1 mm apart from the projecting section  71  of the other primary side core  17 . According to this gap of 1 mm, well convertibility of the primary side cores  17  is achieved. Alternatively, they may be jointed together without any gap. In this case, the leakage flux is avoided between the integral cores  17 . 
     As shown in FIG. 9, the secondary side core  18  is provided around the rotatable gantry section  22  so as to face the primary-side core  17 . As shown in FIG. 10, like the primary side core  17 , the secondary side core  18  is shaped like a squared U. On the central portion of the squared-U shape, the secondary coil  19  is wound. The secondary side core  18  is so arranged that the ends of the squared U are forced to face the static gantry section  12  and conversely the central portion of the squared U is caused to face the rotatable gantry section  22 , thereby making the straight line connecting the two ends of the squared U perpendicular to the direction of rotation of the rotatable gantry section  22 . 
     In FIG. 10, only two cores on the primary side of the high-voltage transformer and only one core on its secondary side are shown. Actually, however, many cores are present on each of the primary side and secondary side and form a circumference as a whole. The many secondary side cores rotate as the rotatable gantry section rotates. The shape, number, and arrangement of the primary and secondary side cores are so determined that the secondary side cores never fail to face the primary side cores. 
     In the present embodiment, the primary side core has two near-rectangular open faces and the secondary side core has two near-square open faces, when viewed from the plane across which the primary and secondary side cores face each other. Making the open faces of the secondary side core smaller than those of the primary side core enables the rotatable gantry section to be made lighter, facilitating insulation. To improve the power supply efficiency, the open faces of the secondary side core may be made larger. Furthermore, the shape of the primary side core and that of the secondary side core may be made the same not only to improve the power supply efficiency but also to facilitate the manufacture. 
     The magnetic flux generated by the current passed through the primary coil  16  reaches the open faces of the secondary side core  18  by way of the open faces (or the ends of the squared U) of the primary side core  17 . The magnetic field generated at that time is shown by a broken-line arrow. The magnetic flux causes the secondary coil  19  wound around the secondary side core  18  to generate current, thereby supplying power from the static gantry section  12  to the rotatable gantry section  22 . 
     With this configuration, power can be supplied continuously, with the primary coil separate from the secondary coil. It should be noted that the flux from the primary side is easily and reliably transmitted to the secondary side according to an existence of the open face (projecting section  71 ). 
     As described above, with the present invention, it is possible to provide a thin-type X-ray computer tomography apparatus with a small-sized high-voltage transformer capable of supplying power and stepping up the voltage to the high voltage which is necessary for generating X-ray at the same time by a noncontacting rotary method. 
     Although no shown, it goes without saying that ripples in the high-voltage output due to the unevenness of the intensity of the magnetic coupling during rotation are removed by the negative feedback of the output voltage, as are ripples resulting from other causes. The negative feedback is effected by optical transmission or by radio. 
     Additional advantages and modifications will readily occur to those skilled in the art. Therefore, the invention in its broader aspects is not limited to the specific details and representative embodiments shown and described herein. Accordingly, various modifications may be made without departing from the spirit or scope of the general inventive concept as defined by the appended claims and their equivalents.

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