Abstract:
A magnetic resonance imaging apparatus generates an MR signal from an object to be examined by applying a gradient field pulse generated by a gradient field coil and a high-frequency magnetic field pulse generated by a high-frequency coil to the object in a static field generated by a static field magnet, and reconstructs an image on the basis of the MR signal. The gradient field coil is housed in a sealed vessel. Numerous techniques are disclosed to reduce adverse effects of vibrations caused by rapidly changing gradient coil currents. By judicious use of non-conducting connection components between gantry components at some joint portions requiring electrical contact and at some other portions not requiring electrical contact, the generation of adverse B waves, and/or induced electron flow in response to physical vibration between joint components can be reduced.

Description:
CROSS REFERENCE TO RELATED APPLICATIONS 
   This is a division of commonly assigned application Ser. No. 09/764,221 filed Jan. 19, 2001 now U.S. Pat. No. 6,556,012 (now allowed). This application is also related to commonly assigned application Ser. No. 09/764,214 filed Jan. 19, 2001 and Ser. No. 09/764,215 filed Jan. 19, 2001. 

   BACKGROUND OF THE INVENTION 
   The present invention relates to a magnetic resonance imaging apparatus for generating a magnetic resonance signal (MR signal) by applying a gradient field pulse and high-frequency magnetic field pulse to an object to be examined which is placed in a homogeneous static field, and generating a magnetic resonance image (MR image) on the basis of this MR signal. 
   In general, a magnetic resonance imaging apparatus of this type includes a static field magnet for generating a static field, a gradient field coil for generating a gradient field, and an RF coil for generating a high-frequency (RF) magnetic field pulse. An object to be examined is placed in the static field formed by the static field magnet. A gradient field pulse and RF magnetic field pulse are applied to this object in accordance with an arbitrarily selected pulse sequence. As a consequence, an MR signal is generated from the object. This MR signal is received via the RF coil. An MR image is reconstructed on the basis of the received MR signal. 
   In the recent technical field of magnetic resonance imaging apparatuses, with improvements in fast imaging technology, research and development have been vigorously carried out. MRI fast imaging requires fast switching of a gradient field and an increase in strength. For this reason, the gradient field coil vibrates. The vibrations cause noise. This noise sometimes reaches 100 db(A) or higher. This makes it necessary for an object to wear earplugs or headphones. 
   A typical measure against noise is to house a gradient field coil in a sealed vessel, as disclosed in Jpn. Pat. Appln. KOKAI Publication No. 63-246146, U.S. Pat. No. 5,793,210, and Jpn. Pat. Appln. KOKAI Publication No. 10-118043. The sealed vessel is evacuated to a nearly vacuum to suppress air-born propagation of noise. 
   Another typical measure against noise is to support a gradient field coil by a damper. This suppresses solid-born propagation of the vibrations of the gradient field coil to the sealed vessel and other parts. 
   Noise can be reduced to a certain degree by such measures against noise. However, it is impossible to further reduce the noise. This is because it is assumed that the gradient field coil is not the only noise source. 
   BRIEF SUMMARY OF THE INVENTION 
   It is an object of the present invention to provide a magnetic resonance imaging apparatus which has an excellent effect of reducing noise and other adverse effects of vibrations. 
   A magnetic resonance imaging apparatus generates an MR signal on an object to be examined by applying a gradient field pulse generated by a gradient field coil and a high-frequency magnetic field pulse generated by a high-frequency coil onto the object in a static field generated by a static field magnet, and reconstructs an image on the basis of the MR signal. The gradient field coil is housed in a sealed vessel. By judicious use of non-conducting connection components between gantry components at some joint portions requiring electrical contact and at some other portions not requiring electrical contact, the generation of adverse B-waves, and/or induced electron flow in response to physical vibration between joint components an be reduced. 
   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 herein after. 

   
     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  is a front view showing the interior of the gantry of a magnetic resonance imaging apparatus according to Embodiment 1 of the present invention; 
       FIG. 2  is a longitudinal sectional view of the gantry according to Embodiment 1; 
       FIG. 3  is an enlarged view of a portion in  FIG. 2 ; 
       FIG. 4  is a perspective view of the gantry according to Embodiment 1; 
       FIG. 5  is a supplementary view of Embodiment 2; 
       FIG. 6  is a schematic view showing an example of the arrangement of a nuclear magnetic resonance imaging apparatus according to Embodiment 2; 
       FIG. 7  is a view showing the arrangement of a whole body RF coil in  FIG. 6 ; 
       FIG. 8  is a developed view of the whole body RF coil in  FIG. 7 ; 
       FIG. 9  is a circuit diagram showing an example of the arrangement of a switch provided for the whole body RF coil; 
       FIG. 10  is an explanatory view showing one connection form of the switches in  FIG. 9  with respect to the whole body RF coil; 
       FIG. 11  is an explanatory view showing a switch connection form different from the one shown in  FIG. 10 ; 
       FIG. 12  is an explanatory view showing a “closed loop” assumed in the whole body RF coil in Embodiment 2; 
       FIG. 13  is an explanatory view showing a switch connection form different from those shown in  FIGS. 10 and 11 ; 
       FIG. 14  is a view for explaining a case where one switch  43   k  is omitted from the switch connection form in  FIG. 13 ; 
       FIG. 15  is an explanatory view showing a switch connection form in which adjacent element loops are left as closed loops in Embodiment 2; 
       FIG. 16  is an explanatory view showing a switch connection form as a modification of the form shown in  FIG. 13 ; 
       FIG. 17  is a perspective view showing an example of the arrangement of an STR type probe in Embodiment 2; 
       FIG. 18  is a schematic view showing another arrangement in Embodiment 2; 
       FIG. 19  is a schematic sectional view of the gantry of a magnetic resonance imaging apparatus according to Embodiment 3; 
       FIG. 20  is a schematic sectional view taken along a plane perpendicular to the Z-axis direction of an ASGC (a plane taken along a line II—II in FIG.  19 ); 
       FIG. 21  is a view for explaining the positional relationship between the shield coil of the ASGC and a side end portion of a static field magnet according to Embodiment 3; 
       FIG. 22  is a view for explaining the compacted winding of the winding portion of the shield coil and a pattern length in Embodiment 3; 
       FIG. 23  is a view for explaining the positional relationship between the shield coil of the ASGC and the side end portion of the static field magnet; 
       FIG. 24A  is a graph showing simulation results on a leak magnetic field and eddy current in a case where the substantial pattern length of the shield coil is equal to the axial length of the magnet; 
       FIG. 24B  is a graph showing simulation results on a leak magnetic field and eddy current in a case where the substantial pattern length of the shield coil is longer than the axial length of the magnet by 10 mm; 
       FIG. 24C  is a graph showing simulation results on a leak magnetic field and eddy current in a case where the substantial pattern length of the shield coil is longer than the axial length of the magnet by 20 mm; 
       FIG. 24D  is a graph showing simulation results on a leak magnetic field and eddy current in a case where the substantial pattern length of the shield coil is longer than the axial length of the magnet by 30 mm; 
       FIG. 25  is a view for explaining the positional relationship between the shield coil of an ASGC and a static field magnet according to Embodiment 3-1; 
       FIG. 26  is a schematic view for explaining the return wire structure of a main coil and shield coil corresponding to each channel in an ASGC according to Embodiment 3-2; 
       FIG. 27  is a view showing the basic arrangement of a magnetic resonance imaging apparatus according to the fourth embodiment; 
       FIG. 28  is a longitudinal sectional view of a gantry according to the fourth embodiment; 
       FIG. 29  is an enlarged view of the portion encircled by the dashed line in  FIG. 28 ; 
       FIG. 30A  is a perspective view of a sealed vessel according to the fourth embodiment; 
       FIG. 30B  is a front view of a sealed vessel according to the fifth embodiment; 
       FIG. 30C  is a partial sectional view of the closed vessel according to the fifth embodiment; 
       FIG. 30D  is a partial sectional view of the closed vessel according to the fifth embodiment; 
       FIG. 31  is a perspective view of a sealed vessel according to the sixth embodiment; 
       FIG. 32  is a cross-sectional view showing how the sealed vessel in  FIG. 31  is joined to a static field magnet vessel; 
       FIG. 33  is a longitudinal sectional view of the cryostat of a static field magnet according to the seventh embodiment; 
       FIG. 34  is a view showing the internal structure of a dynamic vibration absorber in  FIG. 33 ; 
       FIG. 35  is a view showing the internal structure of a cold head portion in another example of the eighth embodiment; 
       FIG. 36  is a longitudinal sectional view of a gantry according to the eighth embodiment; 
       FIG. 37  is a longitudinal sectional view of a gradient field coil unit according to the ninth embodiment; 
       FIG. 38A  is a perspective view showing the principle of the occurrence of noise radio waves in the 10th embodiment; 
       FIG. 38B  is a perspective view showing the principle of the occurrence of noise radio waves in the 10th embodiment; 
       FIG. 39  is a view showing a tuner copper plate and its connection parts in the 10th embodiment; 
       FIG. 40  is a view showing an example of how metal parts are connected to each other in the 10th embodiment; 
       FIG. 41  is a view showing another example of how metal parts are connected to each other in the 10th embodiment; 
       FIG. 42  is a view showing an example of how metal parts are insulated/connected from/to each other in the 10th embodiment; 
       FIG. 43  is a view showing another example of how metal parts are insulated/connected from/to each other the 10th embodiment; 
       FIG. 44  is a view showing still another example of how metal parts are insulated/connected from/to each other the 10th embodiment; 
       FIG. 45  is a perspective view of an RF shield according to the 11th embodiment; 
       FIG. 46  is a longitudinal sectional view of the gantry of a magnetic resonance imaging apparatus according to the 12th embodiment; 
       FIG. 47  is a system diagram of a vacuum pump for a sealed vessel according to the 13th embodiment; 
       FIG. 48  is a graph showing changes in pressure in the sealed vessel in the 13th embodiment; 
       FIG. 49  is a timing chart of ON/OFF operation of the vacuum pump and the opening/closing of valves; 
       FIG. 50  is a view showing the arrangement of the main part of a magnetic resonance imaging apparatus according to the 15th embodiment; 
       FIG. 51  is a view showing the arrangement of the main part of the magnetic resonance imaging apparatus according to the 15th embodiment; 
       FIG. 52A  is a timing chart showing the first driving pattern of a vacuum pump in the 15th embodiment; 
       FIG. 52B  is a timing chart showing the second driving pattern of the vacuum pump in the 15th embodiment; and 
       FIG. 52C  is a timing chart showing the third driving pattern of the vacuum pump in the 15th embodiment. 
   

   DETAILED DESCRIPTION OF THE INVENTION 
   (Embodiment 1) 
     FIG. 1  is a view showing the internal arrangement of the gantry of a magnetic resonance imaging apparatus according to Embodiment 1.  FIG. 2  is a longitudinal sectional view of  FIG. 1. A  substantially cylindrical static field magnet  2001  is comprised of, for example, a superconductive coil that is set in a superconductive state at an extremely low temperature and generates a homogenous static field and a cryostat for maintaining the extremely low temperature state of the superconductive coil. For the sake of convenience, three orthogonal axes (X-, Y-, and Z-axes) are defined with the central axis of the static field magnet  2001  being regarded as the Z-axis. 
   A substantially cylindrical gradient field coil  2002  is placed inside the static field magnet  2001 . The gradient field coil  2002  is an ASGC (Actively Shielded Gradient Coil). The actively shielded gradient coil  2002  is made up of a main coil for generating a gradient field and an active shield coil which is placed outside the main coil to generate a magnetic field in the opposite direction to the gradient field so as to prevent it from leaking out. An RF coil  2010  for transmitting a high-frequency (RF) magnetic field to the object and receiving a magnetic resonance (MR) signal from the object is placed inside the gradient field coil  2002 . 
   The gradient field coil  2002  is supported on an arm  2013  via vibration absorbing members  2012  and position adjustment bolts  2011 . The vibration absorbing member  2012  is made of an elastic material, e.g., rubber, and serves to prevent or suppress the propagation of the solid vibrations of the gradient field coil  2002  to the arm  2013  via the position adjustment bolt  2011 . The arm  2013  is supported on a column  2014  on a base  2015 . 
   The gradient field coil  2002  is housed in a sealed vessel  2003  for an anti-noise measure. The bore of the cryostat of the static field magnet  2001  also serves as the outer wall of the sealed vessel  2003 . The airtightness between the sealed vessel  2003  and the base  2015  is ensured by a metal bellows  2008 . 
   A vacuum pump  2007  is connected to the sealed vessel  2003 . The vacuum pump  2007  evacuates the sealed vessel  2003 . This evacuation maintains a vacuum or a pressure at least lower than the atmospheric pressure in the sealed vessel  2003 . A sound insulating effect is given by
 
 S= 20 log 10 ( P 1/1.01325×10 5 )(decibel:dB)
 
where P1 is the pressure in the sealed vessel  2003 .
 
   If the pressure in the sealed vessel  2003  is, for example, 1,000 pascals, a sound insulating effect of about 40 dB can be obtained. 
   A tube  2018  is connected to the gradient field coil  2002  via the sealed vessel  2003 . Cooling water is circulated via the tube  2018 . 
   The gradient field coil  2002  vibrates due to high-speed switching. The leakage of noise due to this vibration is reduced by the sealed vessel  2003 . Most of solid vibrations are absorbed by the vibration absorbing member  2012 , and the vibrations hardly propagate to the base  2015 . In addition, the following measures are taken against noise and vibrations. 
   The first measure is taken against the vibrations of a cable for sup lying a current from an external power supply to the gradient field coil  2002 . The cable passes through a static field. Owing to high-speed switching, the direction of the current flowing in the cable is inverted at high speed. This current inversion generates the Lorentz force in opposite directions alternately. The cable itself therefore vibrates. 
   The Lorentz force is generated by a current component crossing magnetic lines of force in a static field, The Lorentz force can be reduced by reducing this current component. 
   For this purpose, in this embodiment, as shown in  FIG. 1 , cables  2042  are radially arranged along the radial direction of the cylinder of the static field magnet  2001  to be substantially parallel to the magnetic lines of force in the static field. 
   Each cable  2042  extends from the gradient field coil  2002 , passes through a hole  2043  formed in the sealed vessel  2003 , and is mounted on a cable mount plate  2016  of the cryostat of the static field magnet  2001 . A cable mount portion  2041  of the cable mount plate  2016 , the hole  2043  of the sealed vessel  2003 , and a lead portion  2044  of the cable  2042  from the gradient field coil  2002  are arranged on a straight line passing through the Z-axis. 
   By radially arranging the cables  2042  in this manner, the current component crossing the magnetic lines of force can be reduced to reduce the Lorentz force. This makes it possible to reduce the vibrations of the cables  2042  themselves. 
   The second measure is taken to reduce the solid-born propagation of the vibrations caused by the gradient field coil  2002  and cables  2042  to the sealed vessel  2003  and the cryostat of the static field magnet  2001  via the cables  2042 . 
   The cable  2042  is given flexibility to have a minimum bend radius equal to or less than four times the outer diameter of the cable. This flexibility can be realized by thinning the cable  2042  and selecting a proper coating material. To thin the cable  2042 , a plurality of conductive wires are bundled into a plurality of cables instead of one cable. This can decrease the outer diameter of each cable. In addition, polyethylene or silicone is selected as a coating material for the cable  2042 . 
   An increase in the flexibility of the cable  2042  suppresses the solid-born propagation of the vibrations of the gradient field coil  2002  and static field magnet  2001  to the sealed vessel  2003  and the cryostat of the static field magnet  2001  via the cables  2042 . 
   The third measure is taken for the same purpose as that of the second measure. As described above, the cables  2042  are mounted on the cable mount plate  2016  of the cryostat of the static field magnet  2001 . As shown in  FIG. 3 , the cable mount plate  2016  is attached to an end face of the cryostat of the static field magnet  2001  via the elastic member  2018  such as an antivibration rubber member. With this structure, even if vibrations propagate to the cable mount plate  2016  via the cables  2042 , the vibrations are damped by the elastic member  2018  and hardly propagate to the cryostat of the static field magnet  2001 . 
   The cables  2042  are provided outside via the holes  2043  of the sealed vessel  2003 . Elastic members  2045  as antivibration rubber members are bonded to the holes  2043 . Holes are formed in the elastic members  2045 , and the cables  2042  extend through the holes. With this structure, the vibrations of the cables  2042  are damped by the elastic members  2045  and hardly propagate to the sealed vessel  2003 . 
   The fourth measure is to fix the cable mount plate  2016 ,  2021  on which the cables  2042  are mounted not only to the cryostat of the static field magnet  2001  and but also to the base  2015 . The base  2015  is made of concrete and very heavy and is firmly fixed to a concrete floor with bolts. The base  2015  therefore hardly vibrates. 
   By fixing a cable mount plate  2016 ,  2021  to the base  2015  in this manner, the vibrations of the cable mount plate  2016 ,  2021  can be suppressed. 
   The present invention is not limited to the above embodiment and can be practiced in various modifications. 
   For example, the above four types of measures against noise are independent of each other, and the effect of the present invention can be obtained even if each of the measures is executed singly. In addition, the present invention can be applied to a magnetic resonance imaging apparatus having an apparatus arrangement in which the gradient field coil is not housed in a sealed vessel or an apparatus arrangement in which no vibration absorbing unit is provided for the gradient field coil. Furthermore, the static field generating scheme to be used is not limited to the one using the superconductive coil, and the gradient field coil to be used is not limited to the actively shielded gradient coil. 
   (Embodiment 2) 
   Problems in the prior art which are solved by Embodiment 2 will be described first. 
   As shown in  FIG. 5 , an air-core portion (imaging space) H is formed in a gantry G of an MRI apparatus. A main magnet  3003  for forming a strong static field, a whole body high-frequency coil  3004  for generating an RF magnetic field, and a gradient field coil  3005  are arranged in this air-core portion H, with its central axis L serving as a co-axis. In imaging operation, an object P to be examined is placed on the bed and inserted into the air-core portion H. A compact RF coil is sometimes mounted on the top independently of the whole body high-frequency coil  3004  to observe a small region such as the elbow or knee of the object P. This compact RF coil can be configured for reception only or both transmission and reception. 
   In an MRI apparatus having such an arrangement, de-coupling must be performed between the compact RF coil and the whole body high-frequency coil  3004 . For example, this “de-coupling” is realized by so-called “de-tuning”, i.e., a technique of shifting (de-tuning) the resonance frequency of the compact RF coil from that of the whole body high-frequency coil  3004 . More specifically, such de-tuning is often executed by a technique of short-circuiting the whole body high-frequency coil  3004  to an RF shield  3004   a  mounted outside the coil. In any case, such a treatment is required to prevent a deterioration in detection sensitivity and the like due to electromagnetic coupling or coupling between the whole body high-frequency coil  3004  and the compact RF coil. 
   In some cases, the noise caused by the gradient field coil  3005  is regarded as a problem. This noise is generally the pulse sound generated when the gradient field coil  3005  is driven in accordance with a pulse sequence. In general, this sound may reach 90 dB or 100 dB, which unnecessarily makes the object feel pain or fatigue during an MRI examination. 
   In the prior art, therefore, as shown in  FIG. 5  as well, a means of eliminating the above noise has been proposed, which uses an arrangement in which the gradient field coil  3005  is contained in the region surrounded by an outer wall V and inner wall Va of the sealed vessel. According to such an arrangement, the RF shield  3004   a  is mounted in the sealed vessel, together with the gradient field coil  3005 . This arrangement is used because it is preferable that the whole body high-frequency coil  3004  be located at the greatest distance from the RF shield  3004   a.    
   The following problems, however, arise in the above arrangement. Since the whole body high-frequency coil  3004  and RF shield  3004   a  are disposed outside and inside the sealed vessel, respectively, many electric conductors extending through the inner wall Va of the sealed vessel are required to short-circuit the whole body high-frequency coil  3004  to the RF shield  3004   a  for the purpose of de-tuning. 
   In this arrangement, first, it may be difficult to maintain a predetermined vacuum in the sealed vessel. Second, cumbersome and complicated operations are required to install the MRI apparatus, and maintenance after installation is difficult to perform. 
   Embodiment 2 is configured to solve these problems. The same reference numerals as in  FIG. 5  denote the same parts in the figures to be referred to in the following description. 
     FIG. 6  is a schematic view showing an example of the arrangement of an MRI apparatus according to Embodiment 2. Referring to  FIG. 6 , the MRI apparatus includes a bed  3001 , gantry G, control unit  3006 , and the like. As has been described above, the bed  3001  has a top  3001   a  on which the object P is to be placed. The top  3001   a  can be moved along the body axis of the placed object P. With this movement, the object P can be inserted into the air-core portion (imaging space portion) H in the gantry G. The main magnet  3003 , whole body high-frequency coil  3004 , and gradient field coil  3005  are concentrically arranged around the air-core portion H, with an axis L of the air-core portion H serving as a co-axis. 
   The main magnet  3003  forms a strong static field in the air-core portion H. As this magnet, any of the following magnets can be used: a permanent magnet, electromagnet, superconductive magnet, and the like. 
   The whole body high-frequency coil  3004  is a coil for applying an RF magnetic field (excitation magnetic field) to the object P in the static field to cause nuclear magnetic resonance absorption in the object P. This whole body high-frequency coil  3004  is connected to a transmitter  3004 T via a duplexer  3004 D and driven discretely in terms of time, i.e., on the basis of a pulse signal. A receiver  3021  is connected to the whole body high-frequency coil  3004  via the duplexer  3004 D and receives an NMR signal associated with the object P which is acquired by application of an RF magnetic field. 
   A shown in  FIG. 7 , the whole body high-frequency coil  3004  is formed into a shape similar to a so-called “birdcage” such that a substantially cylindric space is covered with two ring-like conductors  3411  and  3412  that oppose each other and a plurality of linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n , each having two ends connected to a corresponding one of nodes  3411   p1 ,  3411   p2 , . . . ,  3411   pn  and a corresponding one of nodes  3412   p1 ,  3412   p2 , . . . ,  3412   pn  on the peripheral portions of the ring-like conductors  3411  and  3412  . Note that all the ring-like conductors  3411  and  3412  and the plurality of linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n  are formed by electric conductors made of copper or the like, and are “conductive members” of the whole body high-frequency coil  3004 . The above form is generally known as a “birdcage” type. 
   As described above, the ring-like conductors  3411  and  3412  oppose each other, and surfaces  3411 F and  3412 F respectively surrounded by the conductor are parallel to each other. Each of the linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n  is disposed such that it connects a given node (e.g.,  3411   p2 ) on the periphery of the ring-like conductor  3411  to a positionally corresponding one node (e.g.,  3412   p2 ) on the periphery of the ring-like conductor  3412  and is perpendicular to e surfaces  3411 F and  3412 F. These linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n  are therefore parallel to each other. In addition, these linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n , are arranged at predetermined intervals on the peripheries of the ring-like conductors  3411  and  3412 . Note that the respective areas on the side surface of the whole body high-frequency coil  3004  which are defined by the linear conductors  3042   1 ,  3042   2 , . . . ,  3042   n , and ring-like conductors  3411  and  3412  will be referred to as element loops E 1 , E 2 , . . . , En hereinafter. For example, the element loop E 1  is a closed loop surrounded by the four nodes  3411   p1 ,  3412   p1 ,  3412   p2 , and  3411   p2 . 
   A capacitor is inserted in a proper portion of this whole body high-frequency coil  3004 , and resonates with the inductance of a conductive member forming the whole body high-frequency coil  3004 . Various known methods can be used to insert this capacitor. In general, they can be classified into a high-pass type, low-pass type, and bandpass type. The present invention can use any of these forms. 
   The gradient field coil  3005  is a coil for applying different magnetic fields (Gx, Gy, Gz) along three orthogonal axes (x, y, z) defined in the air-core portion H. The degree of gradient of each magnetic field is set by a gradient field power supply system  51 . The positional intersect of an NMR signal received by the whole body high-frequency coil  3004  or a compact RF coil  3002  described below is performed on the basis of the above degree of gradient. 
   In the MRI apparatus of this embodiment, the gradient field coil  3005  is placed in the area surrounded by the outer wall V and inner wall Va of the sealed vessel in the gantry G, as shown in FIG.  6 . This arrangement is configured to prevent noise caused by the gradient field coil  3005  from propagating into the air-core portion. The whole body high-frequency coil  3004  is placed outside the sealed vessel, whereas the RF shield  3004   a  for preventing coupling between the whole body high-frequency coil  3004  and the gradient field coil  3005  is mounted in the sealed vessel. Note that an MRI apparatus having such a sealed vessel will be referred to as a “silent MRI apparatus” hereinafter. 
   This embodiment includes the compact RF coil  3002  which is connected to the receiver  3021 , mounted on the top  3001   a , and used to receive an NMR signal. Although the compact RF coil  3002  is “used to receive”, this coil may be used for transmission by receiving a transmission signal from the transmitter  3004 T via a switch and duplexer (not shown). In some cases, there are merits in terms of imaging operation in letting the compact RF coil  3002  have a transmission function. For example, the transmission energy required by the compact RF coil  3002  located at a knee portion as shown in  FIG. 6  can be reduced by performing transmission from the coil  3002 , and nuclear spins in unnecessary portions are not excited. In a surface coil or the like, the flip angle of an RF pulse changes in accordance with the sensitivity distribution of the coil, resulting in a decrease in NMR signal strength. In general, therefore, the whole body high-frequency coil  3004  is in charge of transmission, and the compact RF coil  3002  is used for reception alone. 
   Referring to  FIG. 6 , the compact RF coil  3002  covers the knee portion of the object P. For example, to observe a small region other than the knee portion, e.g., the elbow of the object P, a coil conforming to the shape of the corresponding small region may be used. 
   According to this embodiment, in the above case, coupling between the two coils  3004  and  3002 , the prevention of which is an object of the present invention, occurs. 
   As shown  FIG. 6 , in addition to the above constituent elements, the MRI apparatus of this embodiment includes, for example, a sequencer  3007  for driving the transmitter  3004 T, a gradient field power supply system  3051 , and the like, the control unit  3006  which performs overall control on the respective constituent elements described above and has an image generating means for reconstructing a tomographic image on the basis of the NMR signal reception result obtained by the whole body high-frequency coil  3004  or compact RF coil  3002 , and a display unit D for displaying the tomographic image and the like. 
   Since the function of the MRI apparatus having the above arrangement is not directly associated with the gist of the present invention and known, a detailed description thereof will be omitted, with only a brief explanation thereof given below. The object P and compact RF coil  3002  on the top  3001   a  are inserted into the air-core portion H of the gantry G, in which a static field is generated by the main magnet  3003 , together with the top  3001   a . The nuclear magnetic moments in the object P are aligned, and Larmor precession of the nuclear magnetic moments is caused along the static field direction as an axis. The RF magnetic field generated by the whole body high-frequency coil  3004  (or compact RF coil  3002 ) is applied to the object to cause magnetic resonance absorption (spin excitation). The resultant NMR signal is then received by the whole body high-frequency coil  3004  or compact RF coil  3002 , and a tomographic image of the object P is reconstructed by the image generating means in the control unit  3006 . In forming the tomographic image, its positional orientation can be performed by the application of a gradient field from the gradient field coil  3005 . 
   In addition, the noise caused when the gradient field coil  3005  is driven (on the basis of a pulse sequence) is reduced by placing the coil  3005  in the sealed vessel. The object P can therefore receive an examination comfortably. 
   The arrangement and function/effect of the whole body high-frequency coil  3004  as a characteristic element of the present invention will be described in further detail below. 
     FIG. 8  is a view obtained by cutting the whole body high-frequency coil  3004  at one point of each of the ring-like conductors  3411  and  3412 , and developing the resultant structure into a plan view. As described with reference to  FIG. 7 , the who e body high-frequency coil  3004  is of the birdcage type constituted by the ring-like conductors  3411  and  3412  and linear conductors  3042   1 ,  3042   2 , . . . ,  3042   12 . Therefore,  FIG. 8 , which is the developed view of the above structure, shows a “ladder-like” structure. In addition, the element loop E 1  on the left end of  FIG. 8  is “identical” to the element loop E 1  on the right end. In addition,  FIG. 8  shows the whole body high-frequency coil  3004  in which the number of element loops is “12” as an example. In general, the birdcage type whole body high-frequency coil  3004  shown in  FIG. 8  is sometimes called a “12-element” whole body high-frequency coil. 
   As shown in  FIG. 8 , in the whole body high-frequency coil  3004 , plurality of switches  3043  are connected in series with the ring-like conductors  3411  and  3412  and linear conductors  3042   1 ,  3042   2 , . . . ,  3042   12 . As a specific form of this switch  3043 , for example, the arrangement shown in  FIG. 9  can be used. Referring to  FIG. 9 , the switch  3043  includes a PIN diode  3431  and inductors  3432 . A control line (not shown) is connected to an end of the inductors  3432 . 
   With these switches  3043 , electromagnetic coupling or coupling between the whole body high-frequency coil  3004  and the compact RF coil  3002  can be disconnected (de-coupling). 
   In the case shown in  FIG. 8 , since the switches  3043  are arranged for the element loops E 2 , E 3 , E 5 , and E 9  and a closed loop “including” the element loops E 2 , E 3 , E 5 , and E 9  (e.g., the closed loop defined by the overall periphery of the element loops E 9  and E 10 ), coupling can be prevented at least for these loops. That is, in general, to prevent coupling between the compact RF coil  3002  and the whole body high-frequency coil  3004 , the switches  3043  are preferably connected to all the loops constituting the whole body high-frequency coil  3004  so as to perform disconnection of the loops. 
   The gist of the present invention is to provide a technique of de-tuning the whole body high-frequency coil  3004  by properly “connecting” or “arranging” the above switches  3043  to the conductor members constituting the whole body high-frequency coil  3004  so as not to cause the above coupling. This “preferred connection form” will be described below. 
   The case shown in  FIG. 10  will be described first as a precondition Referring to  FIG. 10 , the  12  switches  3043  are connected in series with the ring-like conductor  3411 . At first glance, all the element loops E 1 , . . . , E 12  constituting the whole body high-frequency coil  3004  are disconnected in this form. However, a large closed loop corresponding to the ring-like conductor  3412  is left, as indicated by the thick line in FIG.  10 . This is clear from the perspective view of  FIG. 7  instead of the developed view of FIG.  8  . In this case, therefore, effective coupling prevention cannot be expected. That is, for the preferred connection/arrangement form of the switches  3043 , consideration must always be given to the form of the whole body high-frequency coil  3004  as a stereoscopic form. 
   In the form shown in  FIG. 10 , as described above, sufficient decoupling cannot be performed because a closed loop associated with the ring-like conductor  3412  is left. As is obvious from  FIG. 11 , this problem can be solved by connecting or inserting a 13th switch  3043  to or in the ring-like conductor  3412  in addition to the switches  3043  in FIG.  10 . With this arrangement, all the closed loops constituting the whole body high-frequency coil  3004  can be disconnected from each other, and hence perfect decoupling can be executed. This connection/arrangement form of the switches  3043  also falls within the range of the present invention. 
   There is no denying that the arrangement form of the switches  3043  in  FIG. 11  is inefficient. More specifically, with such a means, decoupling can be reliably performed. However, since many switches  3043  must inevitably be prepared, the manufacturing cost increases (the apparatus becomes expensive) accordingly. In addition, the arrangement of the whole body high-frequency coil  3004  itself becomes unnecessarily complicated. 
   With awareness of the problem associated with the form shown in  FIG. 11 , the present inventors propose a more preferable arrangement form of the switches  3043 . To realize “this more preferable arrangement form of the switches 3043”, consider the following general background idea. As the peripheral length or area of a closed loop of various closed loops recognized in the whole body high-frequency coil  3004  in  FIG. 7  or  8  decreases, the degree of coupling to the compact RF coil  3002  decreases. As a consequence, the occurrence of an inconvenience becomes less noticeable. In this case, “various closed loops” include a closed loop defined by the overall periphery of the element loops E 2 , E 3 , and E 4  as indicated by the dashed line in  FIG. 12 , a closed loop including the ring-like conductors  3411  and  3412  as its route as indicated by the chain line in  FIG. 12 , and the like. 
   In consideration of this, if it is checked again whether each of the element loops E 1 , E, . . . , E 12  in the case shown in  FIG. 12  or the like is a “closed loop with the minimum peripheral length”, it is expected that some of these elements need not be disconnected by using the switches  3043 . 
   Referring to  FIG. 13 , on the basis of the above general idea,  30  switches  43  are not provided for the four element loops E 1 , E 4 , E 7 , and E 10  indicated by the thick lines, and the “closed loops” are left established. Although there is a possibility that the above high-frequency current will flow in these element loops E 1 , E 4 , E 7 , and E 10  as well, it can be easily expected that the influence of the above inconvenience is at least much smaller than that in the case where the “closed loop corresponding to the ring-like conductor 3412” shown in  FIG. 10  is present. It was also confirmed that practically sufficient decoupling could be realized when this whole body high-frequency coil  3004  was actually used. 
   In the case shown in  FIG. 13 , a switch  3043   k  provided for the element loop E 5  (or the other switch  3043  provided for the loop E 5 ) plays an important role. More specifically, consider a case where the switch  3043   k  is omitted as shown in FIG.  14 . In this case, the large closed loop indicated by the thick line in  FIG. 14  is left as in the case shown in FIG.  10 . As is apparent from this, two switches  3043  must be connected/arranged to/for at least one of the closed loops with the minimum peripheral length, i.e., the element loops. A large loop in  FIG. 15  is not left. 
   In general, therefore, it is not preferable that two or more element loops, of the element loops E 1 , . . . , E 12 , which have no switches  3043  are continuously present. More specifically, as shown in  FIG. 15 , if no switches  3043  are provided for pairs of adjacent element loops, e.g., the element loops E 2  and E 3 , E 5  and E 6 , E 8  and E 9 , and E 11  and E 12 , and they cannot be disconnected from each other, the closed loops indicated by the thick lines are left. Obviously, the decoupling effect in this case is slightly inferior to that in the case shown in FIG.  13 . Therefore, such a form is preferably avoided. 
   As a modification of the form shown in  FIG. 13 , the connection form of the switches  3043  shown in  FIG. 16  is more preferable. The form shown in  FIG. 16  differs from that shown in  FIG. 13  in that another switch  3043   j  is connected to the element loop E 11 . Although the element loop E 11  has already been disconnected by the presence of a switch  3043   m , the presence of this switch  3043   j  makes it possible to hold symmetry with respect to the switch  3043   k  and reliably obtain a decoupling effect. 
   As described above, according to the whole body high-frequency coil  3004  in this embodiment, a sufficient decoupling effect can be obtained as a whole by determining the connection/arrangement form of the switches  3043  such that at least no large closed loop is left even though some of the element loops E 1 , . . . , E 12  are allowed to have closed loops left as the element loops (minimum peripheral length). This form can decrease the number of switches  3043  to be installed as compared with the form shown in  FIG. 11 , and hence prevents an increase in the manufacturing cost of the whole body high-frequency coil  3004  and complexity of its arrangement. 
   Making the connection form of the switches  3043  in the whole body high-frequency coil  3004  preferable alone contributes to decoupling realized in this embodiment. Therefore, in a silent type MRI apparatus having a sealed vessel like the one shown in  FIG. 6  as well, decoupling between the whole body high-frequency coil  3004  and the compact RF coil  3002  can be realized without posing any problem. 
   Although the above embodiment has exemplified only the “12-element” whole body high-frequency coil, the present invention is not limited to this form. The present invention can be theoretically applied to the whole body high-frequency coil  3004  regardless of the number of elements. 
   The so-called “birdcage type” whole body high-frequency coil  3004  has been described above. However, the present invention is not limited to this form. As is known, various forms have currently been proposed as the forms of whole body high-frequency coils  4 , generally the forms of “NMR probes”. Basically, the present invention can be applied to these forms as long as “decoupling” is required. 
   The various NMR probes described above include a STR (Slotted Tube Rotator) type probe having a cylindrical conductor member like the one shown in FIG.  17  and the like. In this case, switches denoted by reference numeral  3049   k  in  FIG. 17  are preferably connected. In general, decoupling is required for a QD (Quadrature) probe formed by combining a plurality of tuning coils and a multi-surface coil probe, and hence the present invention can be applied to them. 
   In the above embodiment, decoupling between the whole body high-frequency coil  3004  and the compact RF coil  3002  is attained by devising the connection form of the switches  3043  in the whole body high-frequency coil  3004 . This purpose can also be achieved from another viewpoint or by another technique. This technique will be described below as an embodiment different from the above embodiment. 
     FIG. 18  is a schematic view of an MRI apparatus according to another embodiment. Similar to  FIG. 5 ,  FIG. 18  is a front view of a gantry G, which transparently shows part of the interior of the gantry. A characteristic feature of this arrangement in  FIG. 18  is the arrangement form of a whole body high-frequency coil  3004 . More specifically, in the arrangement shown in  FIG. 5 , the whole body high-frequency coil  3004  is placed outside the sealed vessel or on the air-core portion H side with respect to the inner wall Va of the sealed vessel opposing the air-core portion H. The arrangement in  FIG. 18 , however, differs from that in  FIG. 5  in that the whole body high-frequency coil  3004  is also placed in the sealed vessel (the region surrounded by the outer wall V and inner wall Va). 
   In this form, in performing decoupling between the whole body high-frequency coil  3004  and a compact RF coil  3002 , the whole body high-frequency coil  3004  may be de-tuned by short-circuiting the whole body high-frequency coil  3004  to an RF shield  3004   a  as in the technique used in the prior art. In this case, unlike in the arrangement in  FIG. 5 , there is no need to use electric conductors extending trough the wall Va to realize the above short-circuiting. In other words, in the arrangement in  FIG. 18 , since both the whole body high-frequency coil  3004  and the RF shield  3004   a  are arranged in the sealed vessel, the electric conductors connecting them need not extend through the inner wall Va. 
   In this manner, as shown in  FIG. 18 , decoupling between the whole body high-frequency coil  3004  and the compact RF coil  3002  can be achieved. 
   In the form shown in  FIG. 18 , some consideration must be given to the following point. In this case, electric discharge may occur at the two ends of the capacitor forming the whole body high-frequency coil  3004 . To prevent this electric discharge, therefore, a proper molding process is preferably performed for the whole body high-frequency coil  3004  in advance. 
   (Embodiment 3) 
   The problem to be solved by Embodiment 3 will be described first. 
   As a gradient field coil, a shield type coil assembly having a shield coil for suppressing a magnetic field that leaks out is frequently used. As an example of this coil, an ASGC (Actively Shielded Gradient Coil) is known. An ASGC is disclosed in U.S. Pat. Nos. 4,733,189 and 4,737,716. The ASGC has coil assemblies for the generation of magnetic fields corresponding to the X, Y, and Z channels of the MRI apparatus. Each coil assembly has a main coil and shield coil, thus realizing a shield structure for each channel in which a gradient field hardly leaks out. 
   In an MRI apparatus having a conventional shield type gradient field coil, e.g., an ASGC, a leak flux from the ASGC is magnetically coupled to a metal cover located inside a bore forming a diagnosing opening portion (worm bore) and covering a static field magnet to generate an eddy current on the cover. This causes noise. That is, the Lorentz force generated by the eddy current strikes the cover member of the static field magnet to cause noise. 
   For example, in an ASGC, in consideration of the shield performance, a shield coil must ideally (analytically) be wound in nearly infinite pattern length. The “pattern length” is the length of a copper wire wound around a bobbin in a predetermined pattern in the bobbin axial direction (to be referred to as the axial direction herein after), and simply called the axial length in some cases. 
   In practice, the pattern length of the shield coil has its own limit. As a product, in particular, the shorter the pattern length is, the better the product is. On the other hand, a leakage flux from an end portion of the shield coil in the axial direction has no influence on the imaging area spatially formed in the diagnosing open portion. In consideration of the above two states, only the winding pattern of a copper wire on the end portion of the shield coil in the axial direction is packed as much as possible toward the central portion in the axial direction, thus decreasing the pattern length of the overall coil. 
   In such a shield coil, leakage flux is produced from the central portion of the coil pattern in each channel of the ASGC in the axial direction, as a matter of course, regardless of whether the pattern length is decreased or not. This magnet field causes noise as well. When the pattern length of the shield coil is not decreased at the end portion in the axial direction, leakage flux from the ASGC is maximized at the central portion of the pattern. However, as the pattern length is decreased in the above manner, the leakage flux from the end portion of the shield coil in the axial direction increases. That is, when the pattern length is decreased, the leakage flux at the end portion in the axial direction increases, and the noise generated at the end portion of the static field magnet in the axial direction increases. As a consequence, the overall noise increases. 
   According to the prior art, noise generated by a leakage flux at the end portion in the axial direction which does not directly influence the magnetic characteristic of the imaging area, i.e., homogeneity of a magnetic field, has been neglected because of high priority given to a decrease in the axial length of the shield coil. 
   Embodiment 3 has been made in consideration of such situation to maintain the pattern length of a shield coil in the axial direction at the minimum value in an MRI apparatus having an actively (self) shielded gradient coil (ASGC) and suppress noise generated by a leak flux from the end portion in the axial direction. 
   An MRI (Magnetic Resonance Imaging) apparatus according to Embodiment 3 will be described with reference to  FIGS. 19  to  24 . This magnetic resonance imaging apparatus is of a type which includes an actively (self) shielded gradient coil (ASGC). 
     FIG. 19  is a schematic sectional view showing a gantry  4001  of this MRI apparatus along the axial direction. The gantry  4001  has a cylindrical shape as a whole. A worm bore in the gantry serves as a diagnosing space OP. In a diagnosis, an object P to be examined can be inserted into the bore. Note that an X-Y-Z orthogonal coordinate system is set, with the axial direction of the gantry  4001  being defined as the Z-axis. 
   The gantry  4001  includes a static field magnet  4011  which is formed into a substantially cylindrical shape and substantially forms the above bore, a shim coil  4013  attached to, for example, the outer circumferential surface of the gradient field coil  4012 , and an RF coil  4014  placed in the bore of the gradient field coil  4012 . The object P is placed on the top of the bed (not shown) and set in the bore (diagnosing space) while the RF coil  4014  is disposed around the object P. 
   The static field magnet  4011  is formed by an superconductive magnet. That is, a plurality of heat radiation shield vessels and a single liquid helium vessel are housed in an outer vacuum vessel, and a superconductive coil is wound and placed in the liquid helium vessel. The outer circumferential surface of the outer vacuum vessel is covered with a metal cover  4011 A. 
   The ASGC  4012  is formed into an active shield type. This coil  4012  has different coil assemblies for X, Y, and Z channels. Each of the coil assemblies for the respective channels has a shield structure designed to almost prevent a magnetic field from leaking out. In this shield state, the coil generates pulse-like gradient fields in the X-axis, Y-axis, and Z-axis directions. 
   More specifically, as shown in  FIG. 20 , the ASGC  4012  has an X coil assembly  4012 X, Y coil assembly  4012 Y, and Z coil assembly  4012 Z corresponding to the X, Y, and Z channels, which are stacked in units of coils layers in an insulated state to form a substantially cylindrical shape as a whole. Each of the X coil assembly  4012 X, Y coil assembly  4012 Y, and Z coil assembly  4012 Z includes a main coil having a plurality of winding portions for generating a gradient field in a corresponding one of the X-, Y-, and Z-axis directions and a shield coil having a plurality of so-called shield winding portions for suppressing or reducing a leak gradient field (pulse) generated by the winding portions of the main coil. Note that the respective coil assemblies  4012 X,  4012 Y, and  4012 Z are connected to independent gradient field power supplies corresponding to the respective channels. 
     FIG. 21  shows an example of the positional relationship between the main coil and shield coil of the X coil assembly  4012 X or Y coil assembly  4012 Y and a side end face (side end portion)  4011 A a  of the static field magnet  4011 . 
   The Y coil assembly  4012 Y includes four saddle type winding portions (coil patterns) wound around a bobbin B, together with the main coil and shield coil. That is, each coil has two saddle type winding portions juxtaposed in the Z-axis direction and connected in series with each other. Two pairs of such winding portions are arranged to oppose each other in the Y-axis direction. A total of eight winding portions of the main coil and shield coil are electrically connected in series with each other and connected to, for example, a common gradient field power supply. In this case, energization routes are formed such that currents flow in the main coil and shield coil in opposite directions. This makes it possible to generate a gradient field in the Y-axis direction while maintaining the shield function. 
     FIG. 22  shows the coil pattern drawn by one of the four saddle type winding portions of the shield coil. This winding portion CR is actually wound around the bobbin in the form of a saddle, but  FIG. 22  is a plan view of this portion. Each winding portion CR is formed by winding a flat conductor along a substantially spiral pattern in this manner. The winding position of the conductor is analytically obtained on the basis of a predetermined magnetic flux distribution condition. 
   The X coil assembly  4012 X is disposed in the same state as that of the Y coil assembly  4012 Y except that it is rotated about the Z-axis through 90°. 
   The main coil and shield coil of the Z coil assembly  4012 Z are formed by spirally winding flat conductors around the bobbin along analytically obtained winding positions. Each winding portion of the Z coil assembly  4012 Z therefore has a spiral coil pattern. Each of the main coil and shield coil is formed by electrically connecting two winding portions in series, which are wound around the left and right sides of the bobbin in the Z-axis direction. This coil assembly can generate a linear Z-channel gradient field while maintaining a shield function with currents flowing in the respective coils in opposite directions. 
   In each of the coil assemblies  4012 X,  4012 Y, and  4012 Z of the actively shielded gradient coil  4012  having this structure, the winding portion CR of a shield coil  4012   shield  is wound according to a winding method of the present invention. That is, the shield function of the gradient field coil  4012 , i.e., the degree of a leak magnetic field, is determined in association with noise. 
   More specifically, as shown in  FIGS. 22 and 23 , windings L 1  and L 2  are wound around an end portion of the winding portion CR in the Z-axis direction at positions closer to the middle portion in the Z-axis direction than analytically obtained winding positions. The analytically obtained positions of the windings L 1  and L 2  are located outside in the Z-axis direction as indicated by the virtual line in FIG.  23 . By bringing these winding positions closer to the middle portion, the pattern length of the winding portion CR (see  FIG. 22 ) is decreased, which in turn can decrease the required length of the bobbin B in the axial direction. In this case, the desired magnetic field characteristics of the imaging area formed in the diagnosing opening portion are maintained. 
   When the pattern length is decreased in this manner, the magnetic field that leaks out of an end portion of the shield coil  4012   shield  and reaches the metal cover (outer end face)  4011 A placed outside the static field magnet  4011  increases in amount as compared with a case where the pattern length is not decreased. The eddy current generated on the metal cover  4011 A increases due to this leak magnetic field, resulting in an increase in noise due to magnetic coupling. In this embodiment, therefore, a pattern length is set to reduce the amount of noise due to a leak flux from an end portion while maintaining the pattern length of the shield coil  4012   shield  at a small value. 
     FIGS. 24A ,  24 B,  24 C, and  24 D show the simulation results provided for this setting. Each figure shows changes in leak magnetic field Bz and eddy current NI with changes in length D as a parameter which is used to compare the substantial pattern length of the shield coil  4012   shield  in the Z-axis direction and the position of a side end face of the static field magnet  12 . 
   More specifically, referring to  FIGS. 24A ,  24 B,  24 C, and  24 D, in each graph on the left column, the abscissa represents the position of the shield coil  4012   shield  in the Z-axis direction; and the ordinate, the amount of leak magnetic field, whereas in each graph on the right column, the abscissa represents the position of the shield coil  4012   shield  in the Z-axis direction; and the ordinate, the magnitude of eddy current. In both graphs, the abscissa represents the position of the shield coil  4012   shield  in the axial direction on one side (e.g., the positive side) from the center in the Z-axis direction. As both leak magnetic field and eddy current, the values obtained by simulations at predetermined positions on the surface of the metal cover of the gradient field coil  4012  are shown. 
   The “substantial pattern length” in setting the length D indicates a distance to the winding L 1  of the compacted windings L 1  and L 2  forming the shield coil  4012   shield  which is located on the innermost side. 
     FIG. 24A  shows a case where the substantial pattern length of the shield coil  4012   shield  is equal to the axial length of the static field magnet  4011  (i.e., D=0 in FIG.  23 ).  FIG. 24B  shows a case where the substantial pattern length is longer than a side end face of the static field magnet  4011  by 10 [cm] (i.e., D=10 [cm] in FIG.  23 ).  FIG. 24C  shows a case where the substantial pattern length is longer than a side end face of the static field magnet  4011  by 20 [cm] (i.e., D=20 [cm] in FIG.  23 ).  FIG. 24D  shows a case where the substantial pattern length is longer than a side end face of the static field magnet  4011  by 30 [cm] (i.e., D=30 [cm] in FIG.  23 ). 
   As is obvious from these figures, the magnetic field leaking out of the central portion (near the portion represented by Z=30 to 40 [cm]) in the Z-axis direction and the magnitude of eddy current generated by this magnetic field hardly change even with changes in the parameter D representing the positional relationship between the substantial pattern length and the position of the static field end face. In contrast to this, in the end portion (near the portion represented by Z=70 to 80 [cm]) in the axial direction, these amounts greatly change depending on the position relationship parameter D. 
   More specifically, the case shown in  FIG. 24C  (D=20 [cm]) indicates a tendency for the eddy current generated by a leak magnetic field in the end portion (near the portion represented by Z=70 to 80 [cm]) of the static field magnet in the axial direction to become equal or less in magnitude to or than the eddy current generated in the central portion (near the portion represented by Z=30 to 40 [cm]) in the axial direction. That is, if D=20 [cm], the noise produced from the end portion of the static field magnet is equal to or less than the noise produced from the remaining portion of the magnet. Therefore, to suppress noise, the structure represented by D=20 [cm] is optimal. 
   With D=0 or 10 [cm] in  FIG. 24A  or  24 B, the magnitude of eddy current produced in the end portion of the magnet in the axial direction is still larger than that in the central portion. In contrast to this, with D=30 [cm] in  FIG. 24D , although the magnitude of eddy current produced in the end portion of the magnet in the axial direction is smaller than that in the central portion, since the eddy current in the central portion can be regarded as a fixed amount determined by other conditions, it is useless in terms of noise suppression to decrease the magnitude of eddy current produced in the end portion below that in the central portion. 
   The substantial pattern length of the shield coil  4012   shield  of each channel of the ASGC  4012  is therefore maintained at a small value representing a position located outside the side end face of the static field magnet  4011  in the axial direction by distance D=about 20 cm, and noise caused by a leak flux from the end portion can be reduced in amount to that in the central portion. This makes it possible to prevent an increase in size by maintaining the axial length of the shield coil of the ASGC  4012  (i.e., the axial length of the gantry  4001 ) at the minimum value and reduce overall noise as compared with the prior art. 
   (Embodiment 3-1) 
   Embodiment 3-1 of the present invention will be described with reference to FIG.  25 . The same reference numerals as in Embodiment 3 denote the same or equivalent constituent elements in Embodiment 3-1, and a description thereof will be simplified or omitted. A gantry  4001  of Embodiment 3-1 is designed to improve the shape of a side end face of a static field magnet  4011  in consideration of a leak magnetic field from an end portion of a shield coil in the axial direction, i.e., noise due to the resultant eddy current. 
   As shown in  FIG. 25 , a side end face (side end portion)  4011 A a  of the magnet  4011  is rounded and spread from the inner circumferential surface side to the outside in the axial direction, thereby forming a “wide opening shape”. The characteristic feature of this shape can be clearly understood in comparison with the “cylindrical shape” shown in FIG.  21 . 
   With this shape, as the side end face (side end portion)  4011 A a  extends in the axial direction, the distance in the radial direction by which it reaches the metal cover of the magnet increases. As a consequence, the amount of eddy current produced decreases as the end face extends in the axial direction, and hence the amount of noise decreases accordingly. Therefore, the degree of spread of this round portion is determined in consideration of factors such as magnetic field homogeneity in the imaging area formed in a static field and suppression of noise on the basis of simulations or experiments, thereby setting the amount of noise produced from the end portion in the axial direction to be equal to or less than that caused by a leak magnetic field from the remaining portion in the axial direction. In this case, the pattern length of a shield coil  4012   shield  need not be changed from that in the prior art. 
   By forming the side end face (side end portion)  4011 A a  of the static field magnet  4011  into a wide opening shape, noise can be suppressed, and at the same time, the axial length of the ASGC can be set to be shorter than that of the static field magnet, thus making the gantry compact in the axial direction. 
   (Embodiment 3-2) 
   Embodiment 3-2 will be described with reference to FIG.  26 . In each of X, Y, and Z coil assemblies  4012 X,  4012 Y, and  4012 Z of an ASGC  4012  in a gantry  4001  according to Embodiment 3-2, the position of a return wire (also called a connecting wire) which returns from a remote winding portion CR (coil pattern portion) in the radial direction of the bore to the feeding terminal side is determined in consideration of noise suppression. This arrangement may be implemented singly or in combination with the characteristic arrangement of Embodiment 3 or 3-1. 
   In general, the main coil and shield coil corresponding to each channel of an ASGC are wound at different positions in the radial direction of the bore. For this reason, the main coil itself has a return wire, and the shield coil itself has a return wire. That is, since the positions of the return wires of the main coil and shield coil differ from each other in the radial direction of the bore, even if currents flow in the return wires in opposite directions, magnetic fields cannot be canceled out completely. As a consequence, magnetic fields that cannot be canceled out leak from the main coil and shield coil to the outside, causing noise. 
   In this embodiment, therefore, as shown in  FIG. 26 , in the coil assembly corresponding to each channel, a return wire (connecting wire) R shield  of a shield coil  4012   shield  is formed adjacent to a return wire (connecting wire) R main  of a main coil  4012   main  at the same position in the radial direction. In addition, this wiring is performed on the bobbin on the main coil  4012   main  side. The return wire R shield  on the shield coil side may be inserted together with the main coil  4012   main  when it is formed. 
   An up lead L up  extending upward from a feeding terminal T in the radial direction of the bore is connected to the shield coil  4012   shield . A down lead L down  from the shield coil  4012   shield  extends downward at the end portion in the radial direction on the opposite side and is connected to the above return wire R main  to return to another feeding terminal T.  FIG. 26  shows the shape of the winding portion for the Z channel. This return wire structure can equally apply to the X and Y channels. 
   In this structure, currents flow in the two return wires R main  and R shield  in opposite directions. In addition, the positions of the return wires are the same in the radial direction of the bore, and the wires are almost located side by side. As a consequence, most of the magnetic fields produced from these return wires can be canceled out reliably. Therefore, noise due to currents flowing in the return wires can be reliably reduced. In addition, the return wire R shield  on the shield coil side is located on the main coil side, and hence is spaced apart from the static field magnet accordingly. This also contributes to a reduction in noise. Furthermore, in this embodiment, since the pattern length is shorter than the shield coil, a space for routing the return wire R shield  of the shield coil  4012   shield  can be ensured by using a space on the main coil  4012   main  side having a relatively large spatial margin. Since such a space need not be ensured on the shield side, a reduction in the axial size of the ASGC can be ensured. 
   (Fourth Embodiment) 
   The basic arrangement of a magnetic resonance imaging apparatus will be described first with reference to FIG.  27 . The magnetic resonance imaging apparatus includes a gantry  14  having an measurement space in which an object subjected to image diagnosis is to be inserted/placed, a bed  18  disposed adjacent to the gantry  14 , and a control processing section (computer system) for controlling the operations of the gantry  14  and bed  18  and processing MR signals. Typically, a substantially cylindrical measurement space extends through the inner central portion of the gantry  14 . With regard to this cylindrical measurement space, the axial direction is defined as a Z direction, and an X direction (horizontal direction) and Y direction (vertical direction) perpendicular to the Z direction are defined. 
   The gantry  14  has a static field magnet  1  which receives a current supplied from a static field power supply  2  and generates a static field H 0  in the measurement space. This static field magnet  1  is typically formed by a superconductive magnet. The static field magnet  1  has a substantially cylindrical shape as a whole. A gradient field coil  3  is placed in the bore of the static field magnet  1 . The gradient field coil  3  is made up of three coils  3   x ,  3   y , and  3   z  which independently receive currents supplied from a gradient field power supply  4  and generate X-, Y-, and z-axis gradient fields, respectively. The gradient field coil  3  is housed in a sealed vessel in which a vacuum or a similar state is maintained by a vacuum pump. 
   A high-frequency (RF coil)  7  is placed inside the gradient field coil  3 . A transmitter  8 T and receiver  8 R are connected to the RF coil  7 . The transmitter  8 T supplies, to the RF coil  7 , a current pulse that oscillates at a Larmor frequency to excite nuclear magnetic resonance (NMR) under the control of a sequencer  5 . The receiver  8 R receives an MR signal (high-frequency signal) via the RF coil  7 , and performs various kinds of signal processes to form a corresponding digital signal. 
   The sequencer  5  is set under the control of a controller  6  for controlling the overall apparatus. An input device  13  is connected to the controller  6 . The operator can select a desired pulse sequence from a plurality of kinds of pulse sequences in the spin echo method (SE) and echo-planar imaging method (EPI) through the input device  13 . The controller  6  sets the selected pulse sequence in the sequencer  5 . The sequencer  5  controls the application timings of gradient fields in the X-axis, Y-axis, and Z-axis directions, their strengths, the application timing of a high-frequency magnetic field, amplitude, duration, and the like in accordance with the set pulse sequence. 
   An arithmetic unit  10  inputs the MR signal (digital data) formed by the receiver  8 R, and performs processes, e.g., arrangement of measured data in a two-dimensional Fourier space formed in the internal memory and Fourier transform for image reconstruction, to generate image data and spectrum data. A storage unit  11  stores computed image data. A display unit  12  displays an image. 
   An embodiment of the magnetic resonance imaging apparatus having the above basic arrangement will be described next. 
     FIG. 28  is a longitudinal sectional view of the gantry of the magnetic resonance imaging apparatus according to the fourth embodiment. A gradient field coil  102  may be of a non-shield type or active shield type. The gradient field coil  102  has x, y, and z coils as its windings. These x, y, and z coils are housed in a cylindrical bobbin. 
   The gradient field coil  102  having a substantially cylindrical shape is supported on a heavy, concrete gantry base  125  placed on the floor. The gradient field coil  102  is housed in a sealed vessel  133 . The sealed vessel  133  has a liner  131  having a substantially cylindrical shape and forming the inner wall of the vessel, and a vacuum cover  132 . The back surface of the sealed vessel  133  is closed with an inner wall  117  of a cryostat  116  for setting a static field magnet (superconductive coil in this case) in a cryogenic environment. A side wall  118  of the cryostat  116  is joined to the vacuum cover  132  with a joint plate  135 . The sealed vessel  133  is coupled to the gantry base  125  via a vacuum bellows  134  to keep the sealed vessel  133  airtight. 
   The air in the sealed vessel  133  is exhausted by a vacuum pump to keep a vacuum or a similar state in the sealed vessel  133 . This prevents air-born propagation of noise originating from the gradient field coil  102 . 
   An RF coil  103  is placed on the inner surface of the liner  131 . A high-frequency magnetic field is applied to an object via the RF coil  103 , and an MR signal from the object is received. 
   In this arrangement, vacuum leakage tends to occur in the connection portion between the side wall  118  of the cryostat  116  and the joint plate  135 . To prevent this vacuum leakage, an O-ring  108  for vacuum sealing is clamped between the side wall  118  of the cryostat  116  and the joint plate  135 . However, the surface precision of the side wall  118  of the cryostat  116  is not very high. For this reason, the contact precision between the side wall  118  of the cryostat  116  and the O-ring  108  is not very high, and hence the sealing performance of the O-ring  108  is not sufficient. 
   In contrast to this, according to this embodiment, as shown in  FIG. 29 , an annular flange  106  is welded (reference numeral  107 ) to the side wall  118  of the cryostat  116 , and the joint plate  135  of the sealed vessel  133  is fixed to the flange  106  with a bolt  109  via the O-ring  108 . The flange  106  can be formed with high precision by shaving or the like. Since the flange  106  can be brought into contact with the O-ring  108  properly, the sealing performance of the O-ring  108  can be maximized. In addition, since the side wall  118  of the cryostat  116  is connected to the flange  106  by welding, the connection portion therebetween can be kept airtight. This makes it possible to maintain a substantially vacuum state in the sealed vessel  133  and properly prevent air-born propagation of vibrations and noise. 
   (Fifth Embodiment) 
     FIG. 30A  shows an outer appearance of the sealed vessel of a gradient field coil according to the fifth embodiment. To take a measure against noise, the gradient field coil is housed in a sealed vessel  201  held in a substantially vacuum state. In this arrangement, therefore, in the prior art, to check the position of the gradient field coil, the sealed vessel  201  must be partly disassembled. 
   In contrast to this, according to this embodiment, a pair of left and right circular holes are formed in each side wall  207  of the sealed vessel  201 . Windows  202  made of a glass or fiber reinforced plastic material that transmits visible light are fitted in the holes. An operator can easily make a visual check on the position of the gradient field coil in the sealed vessel  201  from the outside via the windows  202 . 
   As shown in  FIG. 30B , a gradient field coil  204  has scale marks  206  each indicating the position of the coil. The scale mark  206  can be visually checked via the window  202 . The operator can objectively grasp the position of the gradient field coil  204  relative to a static field magnet  205  while seeing the scale mark  206 . 
   As shown in  FIG. 30C , leg portions  203  of the sealed vessel  201  have bases  212 . Supports  213  supporting the gradient field coil  204  are fitted in holes vertically formed in the bases  212  to be vertically movable. Threads are formed on the outer surfaces of the supports  213 . Screws  215  are threadably engaged with the threads at crossing axes. When a dial  214  on the distal end portion of each screw  215  is rotated, the support  213  vertically moves, together with the gradient field coil  204 , in the sealed vessel  201 . This makes it possible to adjust the position of the gradient field coil  204  relative to the static field magnet  205 . 
   In this manner, the gradient field coil can be visually checked from the outside without disassembling the vessel, and position adjustment can be performed. This can reduce the chances of degrading airtightness. Therefore, the vessel can be kept airtight, and a sound insulating effect for air-borne propagation of vibrations and noise can be enhanced. 
   Further, as shown in  FIG. 30D , the side walls  207  of the sealed vessel  201  are jointed to the cryostat  217  with the joint plates  235 . Corners where the joint plates  235  are jointed to the side walls  207  are rounded off. Corners where the joint plates  235  are jointed to the liner of the vessel  201  are rounded off. Therefore, the vessel  201  can have a sufficient strength to atmospheric pressure. 
   (Sixth Embodiment) 
     FIG. 31  shows an outer appearance of the sealed vessel of a gradient field coil according to the sixth embodiment. The gradient field coil is housed in a sealed vessel  301 . To prevent air-born propagation of noise originating from the gradient field coil  102 , the air in the sealed vessel  301  is exhausted by a vacuum pump to keep a vacuum or a similar state in the sealed vessel  301 . For this reason, the sealed vessel  133  receives an atmospheric pressure. The strength of the sealed vessel  133  is therefore important. In the fifth embodiment described above, the windows  302  are attached to the side walls  207  of the sealed vessel  201 . In the sixth embodiment, to increase the strength of the portion of each window  302 , a portion of a side wall  304  which surrounds the window  302  is formed into a convex portion  303  having a round shape like a half pipe, thereby reinforcing the portion around the window  302 . 
   With this reinforcement, the degree of vacuum (internal pressure) in the sealed vessel  301  can be sufficiently increased, and hence a sound insulating effect for air-borne propagation of vibrations and noise can be enhanced. 
   As shown in  FIG. 32 , the sealed vessel  301  has a liner  309  having a substantially cylindrical shape and forming the inner wall of the vessel and a vacuum cover  307 . The back surface of the sealed vessel  301  is closed with the inner wall of a cryostat  306  for setting a static field magnet (superconductive coil in this case) in a cryogenic environment. A side wall  311  of the cryostat  306  is joined to the vacuum cover  307 . 
   In an actual manufacturing process, a length L 1  of the cryostat  306  may not match with a length L 2  of an opening portion of the sealed vessel  301  in which the cryostat  306  is to be fitted. In this case, the airtightness of the sealed vessel  301  deteriorates, and vacuum leakage occurs. To solve this problem, in this embodiment, an annular packing  310  is clamped between the liner  309  of the sealed vessel  301  and the vacuum cover  307 . If, therefore, the length L 1  of the cryostat  306  does not match with the length L 2  of the opening portion of the sealed vessel  301  in which the cryostat  306  is to be fitted, the liner  309  of the sealed vessel  301  is joined to the vacuum cover  307  via the packing  310  having a proper width. This makes it possible to easily match the length L 1  of the cryostat  306  with the length L 2  of the opening portion of the sealed vessel  301  in which the cryostat  306  is to be fitted. 
   The packing  310  improves the joining precision between the sealed vessel  301  and the cryostat  306  to improve the airtightness of the sealed vessel  301 . This enhances the sound insulating effect for air-borne propagation of vibrations and noise. 
   (Seventh Embodiment) 
   The gradient field coil is not only a source of vibrations and noise in a magnetic gantry. For example, a heat exchanger using a superconductive coil as a static field magnet produces such vibrations and noise.  FIGS. 33 and 34  are sectional views of a heat exchanger according to this embodiment. A superconductive coil  401  is housed in a cryostat  404 . The cryostat  404  is configured to surround a liquid nitrogen bath housing the superconductive coil  401  together with liquid nitrogen with a plurality of heat radiation shields  402 ,  405 , and  406 . 
   This cryostat  404  has a heat exchanger  407  for absorbing heat from the shield  402  and dissipating it outside. The heat exchanger  407  is comprised of a cylinder  408  having a bottom portion in contact with the shield  402 , a cold head  411  which is cooled by helium gas He and is used to cover the cylinder  408 , a displacer  409  which reciprocates like a piston between the bottom portion and the cold head  411  inside the cylinder  408  with the pressure of helium gas He, and a vacuum bellows  410 . 
   When the displacer  409  is located on the bottom portion, the displacer  409  absorbs heat from the shield  402 . When the displacer  409  is located at the top portion, the displacer  409  transfers heat to the cold head  411 . By repeating this operation, heat can be dissipated from the shield  402 . 
   As described above, since the displacer  409  reciprocates like a piston inside the cylinder  408 , vibrations are produced. The vibrations mechanically propagate to the shields  402 ,  405 , and  406 . This produces noise. 
   To absorb the vibrations, a dynamic vibration absorber  414  is mounted on the cold head  411 . An elastic member, e.g., a spring  412 , of the dynamic vibration absorber  414  is connected onto the cold head  411  such that the expanding direction of the spring  412  is substantially parallel to the direction in which the displacer  409  reciprocates like a piston. A weight  413  is connected to the spring  412 . As the displacer  409  reciprocates like a piston, the weight  413  moves vertically. With this operation, the vibrations of the cold head  411 , originating from the displacer  409 , are absorbed by the dynamic vibration absorber  414 . As a consequence, noise is reduced. 
   The displacer  409  moves like a piston at the frequency of commercial power. The elasticity of the spring  412  and the mass of the weight  413  are set such that the dynamic vibration absorber  414  resonates with vibrations originating from the displacer  409  moving like a piston at this frequency. This makes it possible to effective absorb the vibrations. 
   Vibrations can also be reduced by the following arrangement. As shown in  FIG. 35 , two cylinders  408 - 1  and  408 - 2 , two displacers  409 - 1  and  409 - 2 , and two cold heads  411 - 1  and  411 - 2 , i.e., two heat exchangers, are prepared, and the two heat exchangers are arranged such that the piston motion axes oppose each other, and the displacers  409 - 1  and  409 - 2  are made to move like a piston in opposite phases. 
   (Eighth Embodiment) 
     FIG. 36  is a longitudinal sectional view of the gantry of a magnetic resonance imaging apparatus according to the eighth embodiment. A gradient field coil  502  includes x, y, and z coils as its windings. These x, y, and z coils are housed in a cylindrical bobbin. This substantially cylindrical gradient field coil  502  is supported on a heavy, concrete gantry base  525  installed on the floor. The gradient field coil  502  is housed in a sealed vessel  533 . The sealed vessel  533  includes a liner  531  having a substantially cylindrical shape and forming the inner wall of the vessel and a vacuum cover  532 . The back surface of the sealed vessel  533  is closed with an inner wall  517  of a cryostat  516  for setting a static field magnet (superconductive coil in this case) in a cryogenic environment. A side wall  518  of the cryostat  516  is joined to the vacuum cover  532  with a joint plate  535 . The sealed vessel  533  is coupled to the gantry base  525  with a vacuum bellows  534  to keep the airtightness of the sealed vessel  533 . 
   The vibrations of the gradient field coil  502  mechanically propagate to the sealed vessel  533 . The frequency of the vibrations of the gradient field coil  502  is equal to the alternating frequency of a gradient field in a pulse sequence. Weights  541 ,  542 ,  543 , and  544  are discretely mounted on the liner  531  and vacuum cover  532  such that the liner  531  and vacuum cover  532  of the sealed vessel  533  do not resonate with the vibrations of the gradient field coil  502 , i.e., the natural frequencies of the liner  531  and vacuum cover  532  differ from the vibration frequency of the gradient field coil  502 . 
   The weight  544  mounted on the vacuum cover  532  is, for example, a nonmagnetic metal piece. The annular gel-like substances  541 ,  542 , and  543  are mounted along the inner wall of the liner  531 . The substances  541 ,  542 , and  543  are mounted outside an RF coil  503  to prevent a decrease in the Q value of the RF coil  503 . 
   According to this structure, the liner  531  and vacuum cover  532  of the sealed vessel  533  do not resonate with the vibrations of the gradient field coil  502 . Hence, noise is reduced. 
   Instead of or in addition to mounting the weights on the liner  531  and vacuum cover  532 , the thicknesses of the liner  531  and vacuum cover  532  may be partly decreased. It is an important point of this embodiment that the masses of the liner  531  and vacuum cover  532  are partly increased/decreased to shift their natural frequencies. In addition to shifting the natural frequencies, beams or struts may be used to reinforce the structure. 
   (Ninth Embodiment) 
     FIG. 37  is a longitudinal sectional view of the gantry of a magnetic resonance imaging apparatus according to the ninth embodiment. A gradient field coil  602  includes x, y, and z coils as its windings. These x, y, and z coils are housed in a cylindrical bobbin. This substantially cylindrical gradient field coil  602  is supported on a heavy, concrete gantry base  625  installed on the floor. The gradient field coil  602  is housed in a sealed vessel  633 . The sealed vessel  633  includes a liner  631  having a substantially cylindrical shape, a vacuum cover  532  having a substantially annular, plate-like shape, and a back casing  634  having a substantially cylindrical shape. A cryostat  616  for setting a static field magnet (superconductive coil in this case) in a cryogenic environment is placed outside the back casing  634  of the sealed vessel  633 . An RF coil  635  is mounted on the inner surface of the liner  631 . A high-frequency magnetic field is applied to an object via the RF coil  635 , and an MR signal is received from the object. 
   It is an important point of this embodiment that the sealed vessel  633  housing the gradient field coil  602  does not use the inner wall of the cryostat  616 . In other words, the sealed vessel  633  and cryostat  616  are formed as completely discrete components. If the inner wall of the cryostat  616  is used for the sealed vessel  633  housing the gradient field coil  602 , vacuum leakage tends to occur at the joint portion due to poor surface precision, dimensional errors, and the like of the cryostat  616 . In this embodiment, however, the cryostat  616  is not joined to the sealed vessel  633 . That is, the sealed vessel  633  is manufactured singly. Therefore, high airtightness can be attained regardless of poor surface precision, dimensional errors, and the like of the cryostat  616 . 
   (10th Embodiment) 
   The 10th embodiment is configured to prevent type B waves and induced electrons from being produced when metal parts in the gantry rub against each other, and can be applied to fastening of all metal parts constituting the gantry of a magnetic resonance apparatus which physically vibrates or in which vibrations propagate. 
   The gantry is comprised of many metal parts, which are fastened to each other by mainly using metal screws. If, for example, as shown in  FIG. 38A , when a copper tuner plate  721  is to be mounted on a metal gantry frame  724 , a metal screw  723  and metal insert  722  are generally used in the prior art. Many capacitors are arranged in the gantry. When these capacitors are to be mounted on a tuner plate and the connector of an RF coil tuner is to be fastened to the tuner plate, many metal screws are used. As described above, in the gantry, when parts are to be fixed, metal screws are used at most portions. As shown in  FIG. 38B , when these metal screws rub against the metal parts or metal parts rub against each other due to the above intense vibrations, so-called type B waves are produced. Such type B waves are picked up by the RE coil, and image facts may be produced. This has hardly posed a problem until recently. Recently, however, as higher voltages have been used to attain increases in the speed and strength of a gradient field, type B waves tend to increase in intensity. At present, image artifacts due to increased type B wave noise have become too large to be neglected. In addition to type B waves, electrons induced by contact between, for example, a connector and a tuner plate and vibrations directly enter a signal line to produce image artifacts, problem. 
   It is an object of this embodiment to prevent the occurrence of type B waves and inducted electrons that cause noise. 
   As is known, a gantry is a magnetic apparatus mainly constituted by a static field magnet, gradient field coil, and RF coil, and includes many metal parts. These metal parts are mounted on many portions. These mounting portions can be roughly classified into two types. As shown in  FIGS. 39 and 40 , mounting portions of one type are portions where parts are physically fixed and must be electrically connected to each other, represented by a portion where copper plates constituting an RF coil are attached to each other, a portion where the RF coil copper plates  709  and  710  and a capacitor  711  are attached to each other, a portion where the RF coil copper plate  710  and a lead copper plate  703  are attached to each other, a portion where the lead copper plate  703  and an RF coil tuner copper plate  704  are attached to each other, a portion where the RF coil tuner copper plate  704  and a connector  706  are attached to each other, and a portion where the RF coil tuner copper plate  704  and a capacitor  715  are attached to each other. Mount portions of the other type are portions where it is a main object to physically fix parts to each other, but they need not be electrically connected to each other. 
   It is most preferable that parts be mounted on the former portions by using solder  705 . In this case, since no parts rub against to each other, neither type B wave nor induced electrons are produced. However, solder cannot be used at some portions because of weak fastening force. Screws are used on such portions. 
     FIG. 41  shows an example of how metal parts  731  and  732  are attached to each other by using a resin screw  733 . In the prior art, since a metal screw is used, and the metal screw rubs against the metal parts  731  and  732 , type B waves and induced electrons are inevitably produced. In this embodiment, however, the resin screw  733  is used, and hence generation of such waves and electrons can be prevented. 
     FIG. 42  shows another example of how the metal parts  731  and  732  are attached to each other by using a metal screw  734 . A substantially cylindrical resin spacer  735  is used to prevent direct contact between the metal screw  734  and metal part  731 . In addition, a resin tap  736  is used to prevent contact between the metal screw  734  and the metal part  732 . In this case, although the metal screw  734  is used, type B waves and induced electrons can be prevented by insulating the metal screw  734  from the metal parts  731  and  732  with the resin members  735  and  736 . 
   Obviously, either of the methods shown in  FIGS. 41 and 42  or a combination thereof can be used. It is expected that type B waves and induced electrons will be suppressed by applying the mounting methods shown in  FIGS. 41 and 42  to some portions in the gantry instead of all the corresponding portions. 
   At the portions of the latter type, i.e., the portions where it is the main object to physically fix parts to each other, but there is no need to electrically connect them, metal parts  737  and  738  are attached to each other with the resin screw  733  as shown in, for example, FIG.  43  . In this case, inserting an insulating sheet  739  between the metal parts  737  and  738  can prevent type B waves and induced electrons generated due to friction between the metal parts  737  and  738  as well as type B waves and induced elections generated due to friction between the metal screw and the metal parts as in the prior art. 
     FIG. 44  shows a case wherein the metal parts  737  and  738  are attached to each other by using the metal screw  734 . A substantially cylindrical resin spacer  740  is used to prevent contact between the metal screw  734  and the metal part  738 . In addition, a resin tap  741  is used to prevent contact between the metal screw  734  and the metal part  738 . In this case, although the metal screw  734  is used, the type B waves and inducted electrons can be prevented by insulating the metal screw  734  from the metal parts  737  and  738  with the resin members  740  and  741 . 
   Obviously, either of the methods shown in  FIGS. 43 and 44  or a combination thereof can be used. It is expected that the type B waves and induced electrons will be suppressed by applying the mounting methods shown in  FIGS. 43 and 44  to some portions in the gantry instead of all the corresponding portions. 
   In addition, the type B waves and induced electrons generated due to friction between metal screws and metal parts as in the prior art can be prevented by applying the mounting method shown in  FIG. 43  or  44  to portions where metal parts are attached to resin parts such as a coil bobbin as well as portions where metal parts are attached to each other. 
   (11th Embodiment) 
   The 11th embodiment is related to an improvement in an RF shield placed around an RF coil. The RF shield is typically formed by a copper cylinder to magnetically isolate the RF coil from the outside and shield the RF coil against external electromagnetic noise. An eddy current is produced in this copper cylinder due to high-speed switching of a gradient field, distorting the gradient field. To decrease the time constant of this eddy current, many slits are formed in the copper cylinder. 
   In addition, capacitors are connected between copper plates across the slits to transmit a magnetic field having a relatively low frequency (up to about 100 kHz), e.g., a gradient field, and block a magnetic field having a high frequency of several MHz to several ten MHz, e.g., excitation pulses, i.e., increase a low-frequency impedance and decrease a high-frequency impedance. As another conventional RF shield, an RF shield having capacitances formed on its upper and lower surfaces is also available, which is formed by sticking a plurality of copper plates on the upper and lower surfaces of a dielectric substrate with gaps (slits). 
   A high-speed imaging method such as echo planar imaging (EPI) is required to image, for example, the heart. A very high response speed of a gradient field is indispensable for this operation. For this reason, many slits must be formed in very small increments (at very small intervals). If, however, many slits are formed, the capacitance decreases with a reduction in the area of each copper plate. This makes high-frequency short circuits in the respective slits imperfect. As a consequence, the shield function is made imperfect. 
   This embodiment is configured to achieve both an increase in the number of slits and prevention of a decrease in capacitance. 
     FIG. 45  is a partial perspective view of an RF shield according to this embodiment. A plurality of copper plates  802  are stuck on the upper surface of a dielectric substrate  801  with predetermined gaps (slits)  805 . Likewise, a plurality of copper plates  803  are formed on the lower surface of the dielectric substrate  801  with predetermined gaps (slits)  806 . A capacitance is formed between the copper plates  802  and  803  opposing through the dielectric substrate  801 . 
   In addition, capacitors  804  are formed between the adjacent copper plates  802  on the upper surface of the dielectric substrate  801 . Likewise, capacitors  805  are formed between the adjacent copper plates  803  on the lower surface of the dielectric substrate  801 . 
   In this arrangement, the total capacitance of the capacitors  804  on the upper surface, the capacitors  805  on the lower surface, and the capacitance between the copper plates  802  on the upper surface and the copper plates  803  on the lower surface is ensured as a capacitance large enough to make high-frequency short circuits perfect. 
   (12th Embodiment) 
     FIG. 46  is a longitudinal sectional view of the gantry of a magnetic resonance imaging apparatus according to the 12th embodiment. A gradient field coil  902  includes x, y, and z coils as its windings. These x, y, and z coils are housed in a cylindrical bobbin. This substantially cylindrical gradient field coil  902  is supported on a heavy, concrete gantry base  925  installed on the floor. The gradient field coil  902  is housed in a sealed vessel  933 . The sealed vessel  933  includes a liner  931  having a substantially cylindrical shape and forming the inner wall of the vessel and a vacuum cover  932 . The back surface of the sealed vessel  933  is closed with an inner wall  917  of a cryostat  916  for setting a static field magnet (superconductive coil in this case) in a cryogenic environment. A side wall  918  of the cryostat  916  is joined to the vacuum cover  932  with a joint plate  935 . The sealed vessel  933  is joined to the gantry base  925  with a vacuum bellows  934  to keep the sealed vessel  933  airtight. 
   An RF coil  903  is placed on the inner surface of the liner  931 . A transmitter and receiver are connected to the RF coil  903 . The transmitter supplies a high-frequency current pulse corresponding to a Larmor frequency to the RF coil  903  to excite nuclear magnetization in the object with a high-frequency magnetic field. The transmitter is typically comprised of an oscillating section, phase selecting section, frequency converting section, amplitude modulating section, and high-frequency power amplifying section. The receiver is comprised of a preamplifying section, intermediate frequency converting section, phase detecting section, low-frequency amplifying section, low-pass filter, and A/D converter to receive an MR signal from the object via the RF coil  903 . 
   The transmitter and receiver are housed in an RF unit  940 . The RF unit  940  is installed in a place near the RF coil  903  to achieve reduction in power loss and noise by shortening the cable required. In the prior art, as indicated by the dotted line in  FIG. 31 , the RF unit is mounted on the vacuum cover  932  near an edge portion of an opening portion  941 . In this place, however, the leakage magnetic field from the gradient field coil  902  exhibits the highest strength. The RF unit  940  includes many conductive parts, and eddy currents are produced in these conductive parts due to the leakage magnetic field from the gradient field coil  902 . As a consequence, the conductive parts vibrate due to the Lorents force. The vibrations propagate to the sealed vessel  933  to cause noise. 
   It is an object of this embodiment to reduce noise originating from the RF unit  940 . 
   The RF unit  940  is not mounted on the vacuum cover  932  near the edge portion of the opening portion  941  but is installed in a place physically spaced apart from the sealed vessel  933 , i.e., a place located outside the RF coil  903  at a position near a position directly below the opening portion  941  in the radial direction of the cylindrical gantry with reference to the central axis (Z-axis). More specifically, the RF unit  940  is installed on the heavy, concrete gantry base  925  or another dedicated base. 
   In this installation place, the RF unit  940  is affected less by the leakage magnetic field from the RF coil  903  than in the conventional installation place. For this reason, the vibrations of the conductive parts in the RF unit  940  are reduced. In addition, since the RF coil  903  is physically spaced apart from the sealed vessel  933  and is mounted on the heavy, concrete gantry base  925 , fine vibrations of the RF coil  903  hardly propagate to the sealed vessel  933 . 
   Noise originating from the RF unit  940  can therefore be reduced. 
   (13th Embodiment) 
   As described above, the gradient field is housed in the sealed vessel which is evacuated by the vacuum pump to prevent noise. As the degree of vacuum (pressure) in a sealed vessel increases (decreases), the noise insulating effect increases. To increase the degree of vacuum in the sealed vessel, the vacuum pump is continuously operated during scanning operation in the prior art. This continuous operation shortens the service life of the vacuum pump. If the vacuum pump with decreased capability is used, the degree of vacuum in the sealed vessel cannot be increased, resulting in a deterioration in noise insulating effect. 
   This embodiment is configured to keep a noise insulating effect as long as possible by reducing the load on the vacuum pump. 
     FIG. 47  shows a vacuum pump and piping system according to this embodiment. A sealed vessel  1001  is connected to a vacuum pump  1002  via a main tube  1003 . A solenoid valve  1004  is placed midway along the main tube  1003 . A branch tube  1005  is coupled to the main tube  1003 . The distal end of the branch tube  1005  is open to the atmosphere via a solenoid valve  1006 . 
   The vacuum pump  1002  is driven and the solenoid valves  1004  and  1006  are opened/closed under the control of a pump/valve control section  1020 . The vacuum pump  1002  is alternately driven (ON) and stopped (OFF) under the control of the pump/valve control section  1020 , as shown in FIG.  48 . The duration of an ON period T 1  and the duration of an OFF period T 2  are set in advance such that the pressure in the sealed vessel  1001  does not exceed a predetermined upper limit. The duration of the ON period T 1  and the duration of the OFF period T 2  can be arbitrarily adjusted. 
   Intermittently driving the vacuum pump  1002  in this manner, instead of continuously driving it, can reduce the frequency of maintenance for oil, an oil filter, and the like as compared with a case wherein the vacuum pump  1002  is continuously driven. 
   As shown in  FIG. 49 , the opening/closing of the solenoid valves  1004  and  1006  is interlocked with the intermittent driving of the vacuum pump  1002  by the pump/valve control section  1020 . 
   First of all, the solenoid valve  1006  of the branch tube  1005  is opened/closed in synchronism with the intermittent driving of the vacuum pump  1002 . That is, the solenoid valve  1006  is closed in synchronism with switching of the vacuum pump  1002  from the OFF state to the ON state, and vice versa. 
   To reduce the load on the vacuum pump  1002 , the solenoid valve  1004  of the main tube  1003  is opened with a delay of a time T 3  with respect to the switching timing of the vacuum pump  1002  at which it is switched from the OFF state to the ON state, and is closed a time T 4  earlier than the switching timing of the vacuum pump  1002  at which it is switched from the ON state to the OFF state. These time differences T 3  and T 4  are set to arbitrary times from several sec to several min. 
   Since the solenoid valve  1004  is opened with the delay of the time T 3  from the OFF-to-ON switching timing of the vacuum pump  1002 , lubrication in the vacuum pump  1002  can be completed in a relatively short period of time (pre-vacuum period), i.e., the time T 3 , after the vacuum pump  1002  is started. This is because the object to be evacuated is a small-volume portion extending from the inlet of the pump to the solenoid valve  1004 . When the time T 3  has elapsed after the start of the pump, the solenoid valve  1004  of the main tube  1003  is opened to start evacuating operation (main vacuum) for a target having a large total volume of the volume of the portion extending from the solenoid valve  1004  to the sealed vessel  1001  and the volume of the sealed vessel  1001 . At this time, lubrication in the vacuum pump  1002  has already been completed, and hence the operation can smoothly shift to the main vacuum operation. The load on the vacuum pump  1002  can therefore be reduced. 
   When a predetermined time (T 1  to T 4 ) has elapsed after the vacuum pump  1002  is started, i.e., at a timing the time T 4  earlier than the timing at which the vacuum pump  1002  is turned off, the solenoid valve  1004  of the main tube  1003  is closed. This indicates that the sealed vessel  1001  is isolated from the vacuum pump  1002  when the pressure in the sealed vessel  1001  sufficiently decreases. This makes it possible to prevent an abrupt increase in the pressure in the sealed vessel  1001  upon stopping of the vacuum pump  1002 . 
   (14th Embodiment) 
     FIG. 50  shows the arrangement of the main part of a magnetic resonance imaging apparatus according to the 14th embodiment. A gantry  1101  incorporates a static field magnet  1102  for generating a static field H 0 , a gradient field coil  1103  for receiving a current from a gradient field power supply (G-amp)  1105 , an RF coil  1104 , and a plurality of shim coils  1116  which receive currents from a shim coil power supply (Shim-amp)  1107  and generate magnetic fields for correcting static field inhomogeneity. 
   To achieve noise insulation, the gradient field coil  1103  is housed in a sealed vessel  1115  in which a vacuum or similar state is maintained by a vacuum pump  1111 . A plurality of vacuum sensors (vacuum gages)  1112  are discretely arranged in the sealed vessel  1115  to measure an internal pressure. The data representing the degree of vacuum measured by the vacuum sensor  1112  is stored in a storage section  1113 . Driving state data from the vacuum pump  1111  is stored in the storage section  1113 , together with this degree-of-vacuum data. The driving state data indicates the driving time of the vacuum pump  1111 . 
   A maintenance information generating section  1114  generates maintenance information of the sealed vessel  1115  and vacuum pump  1111  on the basis of the degree-of-vacuum data and driving state data stored in the storage section  1113 , as needed. The maintenance information generating section  1114  generates maintenance information that prompts maintenance of the vacuum pump  1111  and sealed vessel  1115  when it is determined from the degree-of-vacuum data that the degree of vacuum (pressure) in the sealed vessel  1115  does not decrease below a predetermined pressure corresponding to, for example, a noise level of 99 dB in the imaging area. The maintenance information generating section  1114  also generates maintenance information that prompts maintenance of the vacuum pump  1111  when the cumulative driving time calculated from the driving state data exceeds a predetermined value. Each maintenance information is, for example, a message that prompts maintenance of the sealed vessel  1115  or vacuum pump  1111 , and is displayed on a display  1110 . 
   A receiver  1108  acquires an MR signal (high-frequency signal) via the RF coil  1104 , performs pre-processes such as detection and A/D conversion for the signal, and outputs the resultant signal to a processor  1109 . The processor  1109  processes the acquired MR data to generate an image and spectrum. The image and spectrum are sent to the display  1110  to be displayed. 
   The processor  1109  has the function of correcting the phase of the MR data acquired by the receiver  1108  and performing frequency shift on the basis of degree-of-vacuum data as well as the main function of generating images and spectra. As the degree of vacuum varies, the strength H 0  of the static field varies. As the strength H 0  of the static field varies, for example, a resonance frequency f 0  of a proton varies in the static field on which no gradient field is superimposed. The processor  1109  holds data representing the relationship between the degree of vacuum measured in advance and the resonance frequency f 0 , and specifies the resonance frequency (corrected resonance frequency) f 0  corresponding to the degree-of-vacuum data by referring to the relationship data. In MRS (MR spectroscopy), the phase of the MR data acquired by the receiver  1108  is corrected and frequency shift is performed on the basis of this corrected resonance frequency f 0 . The processor  1109  then generates a spectrum on the basis of this corrected data. In practice, data is repeatedly acquired, and phase correction and frequency shift are performed for each data to generates a plurality of spectra. These spectra are then added together. In EPI (Echo Planar Imaging), an EPI image is generated on the basis of acquired data, and the EPI image is shifted in the phase encoding direction (the shifting of the EPI image largely generates in the phase-encoding direction, and generates in the read-out direction in a small). In practice, data is repeatedly acquired, and an EPI image is generated for each data. Each image is then shifted in the phase encoding direction, and the resultant EPI images are added/subtracted. In the case of a phase image as well, a phase shift amount is calculated on the basis of the corrected resonance frequency f 0 , and the phase image is corrected on the basis of the phase shift amount. 
   As described above, according to this embodiment, maintenance information can be generated, as needed. In addition, phase and frequency correction can be made in accordance with variations in degree of vacuum. 
   (15th Embodiment) 
     FIG. 51  shows the arrangement of the main part of a magnetic resonance imaging apparatus according to the 15th embodiment. A gantry  1201  incorporates a static field magnet  1202  for generating a static field H 0 , a gradient field coil  1203  for receiving a current from a gradient field power supply (G-amp)  1205 , an RF coil  1204  connected to a transmitter/receiver (RF-amp)  1208 , and a plurality of shim coils  1216  which receive currents from a shim coil power supply (Shim-amp)  1207  and generate magnetic fields for correcting static field inhomogeneity. 
   To achieve noise insulation, the gradient field coil  1203  is housed in a sealed vessel  1215  in which a vacuum or similar state is maintained by a vacuum pump  1211 . A plurality of vacuum sensors (vacuum gages)  1212  are discretely arranged in the sealed vessel  1215  to measure an internal pressure. On the basis of the degree-of-vacuum data measured by the vacuum sensor  1212 , a real-time manager  1210  outputs an instruction, e.g., an instruction to wait for the execution of a pulse sequence to a sequence controller  1209  for controlling the gradient field power supply  1205 , transmitter/receiver  1208 , and shim coil power supply  1207  in accordance with the pulse sequence. The real-time manager  1210  also controls the operation of the vacuum pump  1211  on the basis of the measured degree-of-vacuum data. Note that a system manager  1213  is used to control the overall system in accordance with an instruction input by an operator through a console  1214 . 
   Real-time control of the real-time manager  1210  will be described first. The real-time manager  1210  executes the following functions. 
   (1) The vacuum pump  1211  is started before scanning operation. The real-time manager  1210  does not output a scan start command to the sequence controller  1209  until the degree of vacuum in the sealed vessel  1215  (pressure in the sealed vessel) decreases below a predetermined value. That is, the real-time manager  1210  outputs a scan start command to the sequence controller  1209  only when the degree of vacuum exceeds the predetermined value. 
   (2) In executing a pulse sequence sensitive to magnetic field variations, e.g., MRS or EPI, the real-time manager  1210  continuously drives the vacuum pump  1211  during scanning operation. 
   (3) When the degree of vacuum exceeds the predetermined value during scanning operation, the real-time manager  1210  outputs a command to stop the scanning operation to the sequence controller  1209 . 
   (4) When the degree of vacuum decreases below the predetermined value, the real-time manager  1210  drives the vacuum pump  1211  before scanning operation, and does not output a scan start command to the sequence controller  1209  until the degree of vacuum reaches a predetermined value. 
   (5) The real-time manager  1210  selectively uses a driving pattern for the vacuum pump  1211  in accordance with imaging conditions (e.g., the type of pulse sequence, an average number, and dynamic imaging). In executing, for example, a pulse sequence in the spin echo method or the like, which is not very sensitive to magnetic field variations, the real-time manager  1210  intermittently drives the vacuum pump  1211 , as shown in FIG.  52 A. For example, the real-time manager  1210  drives the vacuum pump  1211  for a period ΔT 1 , and stops it for a period Δt 1 . The vacuum pump  1211  is alternately driven/stopped repeatedly in this manner. In executing a pulse sequence which is relatively sensitive to magnetic field variations, as shown in  FIG. 52B , the real-time manager  1210  sets a driving period ΔT 2  and stop period Δt 2  of the pump  1211  to be shorter than ΔT 1  and Δt 1 , thus reducing the width of magnetic field variations. In executing a pulse sequence which is very sensitive to magnetic field variations, e.g., MRS or EPI, the real-time manager  1210  continuously drives the vacuum pump  1211  as shown in  FIG. 52C  in the same manner as in (2). In addition, in executing a pulse sequence which is very sensitive to magnetic field variations, e.g., MRS or EPI, the real-time manager  1210  may stop the pump  1211  and set the atmospheric pressure in the sealed vessel instead of continuously driving the pump  1211 . In this case, although a noise reducing effect cannot be expected, at least magnetic field variations can be eliminated. To properly reconstruct images even at the atmospheric pressure, the real-time manager  1210  holds image quality parameter (magnetic field inhomogeneity, center frequency, and phase shift) information corresponding to the atmospheric pressure in advance, and the transmitter/receiver  1208  adjusts the shim coil current, the center frequency and phase of a high-frequency current pulse in the transmitter/receiver  1208 , and the reference frequency and phase of a reception system in accordance with these parameters. 
   (6) The real-time manager  1210  drives/stops the pump  1211  in accordance with the comparison result between the measured degree of vacuum and the predetermined value. More specifically, when the measured degree of vacuum exceeds an upper limit, the real-time manager  1210  drives the pump  1211 . When the measured degree of vacuum is below a lower limit, the real-time manager  1210  stops the pump  1211 . This makes it possible to suppress variations in degree of vacuum between the upper limit value and the lower limit value. The upper and lower limit values can be changed in accordance with imaging conditions as in the case of (5). 
   (7) If the degree of vacuum does not decrease below the predetermined value even after the pump  1211  is continuously driven, a warning is generated by sound or image display. 
   The real-time manager  1210  also has the function of performing the following corrections in accordance with the degree of vacuum. (1) Magnetic field inhomogeneity changes depending on the degree of vacuum. The relationship between the degree of vacuum and magnetic field inhomogeneity is measured and held in the real-time manager  1210  in advance. The real-time manager  1210  specifies magnetic field inhomogeneity in accordance with the degree of vacuum by referring to this relationship, and adjusts the shim coil current to be supplied to the shim coil power supply  1207  in accordance with the specified magnetic field inhomogeneity. This makes it possible to quickly correct magnetic field inhomogeneity. In practice, the relationship between the degree of vacuum and magnetic field inhomogeneity is discretely measured, and magnetic field inhomogeneity can be obtained by linear interpolation from the discrete value. (2) As the degree of vacuum varies, the strength of the static field varies. As a result, a resonance frequency B 0  of a proton varies in the static field on which no gradient field is superimposed. The real-time manager  1210  adjusts the center frequency and phase of a high-frequency current pulse in the transmission system of the transmitter/receiver  1208  in accordance with the resonance frequency B 0  corresponding to this degree of vacuum. In addition, the real-time manager  1210  adjusts the reference frequency and phase of the reception system. 
   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.