Source: http://www.google.ca/patents/US7071693
Timestamp: 2015-08-01 03:36:34
Document Index: 77184203

Matched Legal Cases: ['arts 731', 'arts 731', 'arts 731', 'art 731', 'art 732', 'arts 731', 'arts 737', 'arts 737', 'arts 737', 'arts 737', 'art 738', 'art 738', 'arts 737']

Patent US7071693 - Magnetic resonance imaging apparatus - Google PatentsSearch Images Maps Play YouTube News Gmail Drive More »Sign inAdvanced Patent SearchPatentsA 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,...http://www.google.ca/patents/US7071693?utm_source=gb-gplus-sharePatent US7071693 - Magnetic resonance imaging apparatusAdvanced Patent SearchPublication numberUS7071693 B2Publication typeGrantApplication numberUS 10/303,720Publication date4 Jul 2006Filing date26 Nov 2002Priority date21 Jan 2000Fee statusPaidAlso published asUS6556012, US20010022515, US20030107376Publication number10303720, 303720, US 7071693 B2, US 7071693B2, US-B2-7071693, US7071693 B2, US7071693B2InventorsYasutake YasuharaOriginal AssigneeKabushiki Kaisha ToshibaExport CitationBiBTeX, EndNote, RefManPatent Citations (39), Non-Patent Citations (1), Referenced by (16), Classifications (6), Legal Events (3) External Links: USPTO, USPTO Assignment, EspacenetMagnetic resonance imaging apparatus
US 7071693 B2Abstract
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.
1. A magnetic resonance imaging apparatus comprising a gantry whose structural components include: a static field magnet, a gradient field coil, a high-frequency coil, and a sealed vessel that houses the gradient field coil, and
wherein at least one of the structural components of the gantry which is formed by one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel is coupled together with another one of the structural components of the gantry which is formed by another one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel at least partially with attachment screws such that: (a) where electric connection is require, coupled structural components of the gantry are in metallic contact with one another, and attachment screws are prevented from being in metallic contact with the joined structural components, and (b) where physical fixing is required but electrical connection is not required, coupled structural components of the gantry and attachment screws are prevented from being in mutual metallic contact with one another. 2. A magnetic resonance imaging apparatus as in claim 1 wherein said attachment screw comprise resin screws.
3. A magnetic resonance imaging apparatus as in claim 1, wherein said attachment screws comprise metallic screws, and said metallic screws are covered with resin to prevent the metallic screws from being in contact with the thereby joined structural components.
4. A magnetic resonance imaging apparatus as in claim 1 wherein structural components are prevented from being in mutual metallic contact with one another by use of an insulating sheet.
5. A magnetic resonance imaging apparatus as in claim 1 wherein structural components are formed of a metallic material.
6. A magnetic resonance imaging apparatus as in claim 2 wherein structural components are formed of a metallic material.
7. A magnetic resonance imaging apparatus as in claim 3 wherein structural components are formed of a metallic material.
8. A magnetic resonance imaging apparatus as in claim 4 wherein structural components are formed of a metallic material.
9. A magnetic resonance imaging apparatus comprising a gantry whose structural components include a static field magnet, a gradient field coil, a high-frequency coil, and a sealed vessel that houses the gradient field coil, and
wherein at least one of the structural components of the gantry which is formed by one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel is coupled together with another one of the structural components of the gantry which is formed by another one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel such that, at where electric connection is required, the connected gantry structural components are prevented from being in metallic contact with attachment screws used for joining such components into electrical connection. 10. A method for joining structure components of an MRI gantry whose structural components include: a static field magnet, a gradient field coil, a high-frequency coil, and a sealed vessel that houses the gradient field coil, said method comprising:
coupling at least one of the structural components of the gantry which is formed by one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel with another one of the structural components of the gantry which is formed by another one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel together at least partially with attachment screws such that: (a) where electric connection is required structural components of the gantry are in metallic contact with one other, and attachment screws are prevented from being in metallic contact with the joined structural components, and (b) where physical fixing is required but electrical connection is not required, structural components of the gantry and attachment screws are prevented from being in mutual metallic contact with one another. 11. A method as in claim 10 wherein said attachment screw comprise resin screws.
12. A method as in claim 10 wherein said attachment screws comprise metallic screws, and said metallic screws are covered with resin to prevent the metallic screws from being in contact with the thereby joined structural components.
13. A method as in claim 10 wherein structural components are prevented from being in mutual metallic contact with on another by use of an insulating sheet.
14. A method as in claim 10 wherein structural components are formed of a metallic material.
15. A method as in claim 11 wherein structural components are formed of a metallic material.
16. A method as in claim 12 wherein structural components are formed of a metallic material.
17. A method as in claim 13 wherein structural components are formed of a metallic material.
18. A method for joining structural components of an MRI gantry whose structural components include a static field magnet, a gradient field coil, a high-frequency coil, and a sealed vessel that houses the gradient field coil, said method comprising:
coupling at least one of the structural components of the gantry which is formed by one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel with another one of the structural components of the gantry which is formed by another one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel such that, where electric connection is required, the connected gantry structural components are prevented from being in metallic contact with attachment screws used for joining such components into electrical connection. 19. A method for physically connecting MRI gantry components together in electrical contact while reducing generation of adverse electrical signals otherwise caused by physical vibrations between such joined gantry components, said method comprising:
clamping two MRI gantry components together in electrical contact using a non-metallic fastening structure in contact with said components. 20. A method as in claim 19 wherein said non-metallic fastening structure comprises a non-metallic screw.
21. A method as in claim 19 wherein said non-metallic fastening structure comprises a metallic screw in contact with a least one noted non-metallic member which member is, in turn, in contact with at least one of the joined gantry components.
22. A method as in claim 19 further comprising:
clamping at least two MRI gantry components together without permitting electrical contact using at least one non-metallic fastening structure to clamp the components together physically without making electrical contact therebetween. 23. A method as in claim 22 wherein said non-metallic fastening structure comprises a non-metallic screw.
24. A method as in claim 22 wherein said non-metallic fastening structure comprises a metallic screw in contact with at least one noted non-metallic member which member is, in turn, in contact with at least one of the joined gantry components.
25. An MRI gantry including a static field magnet, a gradient field coil, a high-frequency coil, and a sealed vessel that house the gradient field coil, said gantry comprising:
at least one of the structural components which is formed by one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel is connected together with another one of the structural components of the gantry which is formed by another one of the static field magnet, gradient field coil, high-frequency coil and sealed vessel where electric connection is required but which are prevented from being in metallic contact with an attachment member used for joining such components into electrical connection. 26. An MRI gantry having components connected together in electrical contact while reducing generation of adverse electrical signals otherwise caused by physical vibrations between such joined gantry components, said gantry comprising:
two MRI gantry components clamped together in electrical contact using a non-metallic fastening structure in contact with said components. 27. An MRI gantry as in claim 26 wherein said non-metallic fastening structure comprises a non-metallic screw.
28. An MRI gantry as in claim 26 wherein said non-metallic fastening structure comprises a metallic screw in contact with at least one noted non-metallic member which member is, in turn, in contact with at least one of the joined gantry components.
29. An MRI gantry as in claim 26 further comprising:
at least two further MRI gantry components that are clamped together without permitting electrical contact using at least one non-metallic fastening structure to clamp the components together physically without making electrical contact therebetween. 30. An MRI gantry as in claim 29 wherein said non-metallic fastening structure comprises a non-metallic screw.
31. An MRI gantry as in claim 29 wherein said non-metallic fastening structure comprises a metallic screw in contact with at least one noted non-metallic member which member is, in turn, in contact with at least one of the joined gantry components.
32. A magnetic resonance imaging apparatus comprising a gantry whose structural components include a radio frequency (RF) coil having a metallic plate,
wherein the metallic plate of the RF coil is physically fixed and electrically connected to another component of the gantry by an attachment screw which is prevented from being in mutual metallic contact with the metallic plate of the RF coil and the other component of the gantry. 33. A magnetic imaging apparatus as in claim 32, wherein the other component of the gantry is another metallic plate of the RF coil.
34. A magnetic imaging apparatus as in claim 32, wherein the other component of the gantry is a capacitor.
35. A magnetic imaging apparatus as in claim 32, wherein the other component of the gantry is a lead copper plate.
36. A magnetic imaging apparatus as in claim 32, wherein the other component of the gantry is a connector.
37. A method of joining a metallic plate of a radio frequency (RF) coil of an MRI gantry to another component of the MRI gantry so that the metallic plate of the RF coil and the other component of the MRI gantry is physically fixed and electrically connected to each other by an attachment screw which is prevented from being in metallic contact with both the metallic plate of the RF coil and the other MRI gantry component.
38. A method for physically and electrically connecting an metallic plate of a radio frequency (RF) coil of an MRI gantry to another component of the MRI gantry, the method comprising clamping the metallic plate of the RF coil together with the other component of the MRI gantry using a non-metallic fastening structure in contact with the metallic plate of the RF coil and the other component of the MRI gantry.
39. A method as in claim 38, wherein said non-metallic fastening structure comprises a non-metallic screw.
40. A method as in claim 38, wherein said non-metallic fastening structure comprises a metallic screw in contact with a non-metallic member which is, in turn, in contact with at least one of the metallic plate of the RF coil and the other MRI gantry component.
41. A method as in claim 38, wherein the other component of the MRI gantry is another metallic plate of the RF coil.
42. A method as in claim 38, wherein the other component of the MRI gantry is a capacitor.
43. A method as in claim 38, wherein the other component of the MRI gantry is a lead copper plate.
44. A method as in claim 38, wherein the other component of the MRI gantry is a connector.
45. An MRI gantry comprising a radio frequency (RF) oil, the RF coil having metallic plates which are physically and electrically connected together by an attachment member which is prevented from being in metallic contact with the metallic plates.
46. A method as in claim 45, wherein the attachment member is a non-metallic screw.
47. A method as in claim 45, wherein the attachment member is a metallic screw in contact with a non-metallic member which is, in turn, in contact with at least one of the metallic plates of the RF coil.
48. An MRI gantry comprising a radio frequency (RF) oil having a metallic tuner plate which is physically and electrically connected to another MRI gantry component by an attachment member, the attachment member prevented from being in metallic contact with the metallic tuner plate and the another MRI gantry component.
49. An MRI gantry as in claim 48, wherein the another MRI gantry component is a connector.
50. An MRI gantry as in claim 48, wherein the another MRI gantry component is a capacitor.
51. An MRI gantry as in claim 48, wherein the attachment member is a non-metallic screw.
52. An MRI gantry as in claim 48, wherein the attachment member is a metallic screw in contact with a non-metallic member which is, in turn, in contact with at least one of the metallic RF tuner plate and the another MRI gantry component.
53. A method of physically and electrically connecting metallic tuner plate of a radio frequency (RF) coil of an MRI gantry with another MRI gantry component, the method comprising physically and electrically connecting the metallic tuner plate of the RF coil to the another MRI gantry component by an attachment member, the attachment member prevented from being in metallic contact with the metallic tuner plate of the RF coil and the another MRI gantry component.
54. A method as in claim 53, wherein the another MRI gantry component is a connector.
55. A method as in claim 53, wherein the another MRI gantry component is a capacitor.
56. A method as in claim 53, wherein the attachment member is a non-metallic screw.
57. A method as in claim 54, wherein the attachment member is a metallic screw in contact with a non-metallic member which is, in turn, in contact with at least one of the metallic RF tuner plate and the another MRI gantry component.
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.
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.
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.
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.
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 log10(P1/1.01325�105)(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.
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 3004T via a duplexer 3004D 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 3004D 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 3411F and 3412F 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 3411F and 3412F. 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 E1, E2, . . . , En hereinafter. For example, the element loop E1 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 3004T 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 3004T, 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 E1 on the left end of FIG. 8 is “identical” to the element loop E1 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 E2, E3, E5, and E9 and a closed loop “including” the element loops E2, E3, E5, and E9 (e.g., the closed loop defined by the overall periphery of the element loops E9 and E10), 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 E1, . . . , E12 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 E2, E3, and E4 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 E1, E, . . . , E12 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 E1, E4, E7, and E10 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 E1, E4, E7, and E10 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 E5 (or the other switch 3043 provided for the loop E5) 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 E1, . . . , E12, 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 E2 and E3, E5 and E6, E8 and E9, and E11 and E12, 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 E11. Although the element loop E11 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 E1, . . . , E12 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.
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 4011A.
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 4012X, Y coil assembly 4012Y, and Z coil assembly 4012Z 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 4012X, Y coil assembly 4012Y, and Z coil assembly 4012Z 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 4012X, 4012Y, and 4012Z 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 4012X or Y coil assembly 4012Y and a side end face (side end portion) 4011Aa of the static field magnet 4011.
The Y coil assembly 4012Y 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 4012X is disposed in the same state as that of the Y coil assembly 4012Y except that it is rotated about the Z-axis through 90�.
The main coil and shield coil of the Z coil assembly 4012Z are formed by spirally winding flat conductors around the bobbin along analytically obtained winding positions. Each winding portion of the Z coil assembly 4012Z 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 4012X, 4012Y, and 4012Z 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 L1 and L2 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 L1 and L2 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) 4011A 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 4011A 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, 24B, 24C, and 24D 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, 24B, 24C, and 24D, 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 L1 of the compacted windings L1 and L2 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 24B, 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 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) 4011Aa 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) 4011Aa 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) 4011Aa 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 will be described with reference to FIG. 26. In each of X, Y, and Z coil assemblies 4012X, 4012Y, and 4012Z 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) Rshield of a shield coil 4012 shield is formed adjacent to a return wire (connecting wire) Rmain 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 Rshield on the shield coil side may be inserted together with the main coil 4012 main when it is formed.
An up lead Lup 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 Ldown 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 Rmain 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 Rmain and Rshield 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 Rshield 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 Rshield 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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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.
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 T1 and the duration of an OFF period T2 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 T1 and the duration of the OFF period T2 can be arbitrarily adjusted.
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.
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 H0, 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.
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 H0, 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.
(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. 52A. For example, the real-time manager 1210 drives the vacuum pump 1211 for a period ΔT1, and stops it for a period Δt1. 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 ΔT2 and stop period Δt2 of the pump 1211 to be shorter than ΔT1 and Δt1, 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.
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