An MRI apparatus measures a signal obtained from a nuclear magnetic resonance (hereinafter referred to as “NMR”) phenomenon of an examinee disposed under a static magnetic field, and subjects the signal to calculation processing to display images of a density distribution, a relaxation time distribution, etc. of nuclear spins in the examinee as tomograms, and it is used for various kinds of diagnosis, etc. while targeting human bodies as examinees.
In the MRI apparatus, an examinee is disposed in a static magnetic field space (imaging space) having spatially and timely uniform intensity and direction which is formed by a magnet, pulsed electromagnetic waves are applied to the examinee by an RF coil, and a generated NMR signal is received by the RF coil. At the reception time, in order to add position information to the NMR signal, a gradient magnetic field is generated in the imaging space by a triaxially-orthogonal gradient magnetic field coil.
An active shield type gradient magnetic field coil (Active Shield Gradient Coil: ASGC) can be adopted as the gradient magnetic field coil. ASGC has a main coil and a shield coil in each of three axial directions, so that the pulse magnetic field leaking to the magnet side can be reduced. Therefore, it is possible to reduce unnecessary magnetic field which is transiently varied by eddy current adversely affecting MRI imaging.
In the MRI apparatus as described above, the total thickness of the elements from the RF coil till the gradient magnetic field coil, containing a cover, is typically equal to about 150 mm. In order to secure a broad imaging space in which an examinee is disposed, it is considered that the internal diameter of a magnet is increased by the amount corresponding to the total thickness from the thickness of the cover till the thickness of the gradient magnetic field coil. However, in this case, the static magnetic field generating efficiency is lowered and thus large magnetomotive force is needed. For example, in the case of a superconducting magnet, many superconducting wires are used. Therefore, the cost of a magnet is increased. Conversely, it is also considered that the internal diameter of the magnet is not varied and the thickness of the gradient magnetic field coil is reduced. However, in this case, the interval between the main coil and the shield coil is reduced, so that the generation efficiency of the gradient magnetic field is lowered and thus a lot of current is needed to generate a desired gradient magnetic field.
Therefore, for example, Patent Document 1 has proposed the structure that a recess portion is provided at the center portion of a gradient magnetic field coil and an RF coil is disposed at the recess portion. By this structure, a large imaging space can be secured, and the interval between the main coil and the shield coil can be increased at the end portion area of the gradient magnetic field coil. Therefore, a desired gradient magnetic field can be generated in the imaging space by small current, and thus the generation efficiency of the gradient magnetic field is excellent.
Patent Document 1: JP-T-2005-515051
However, the structure that the recess portion is provided at the center portion of the gradient magnetic field coil as in the case of the Patent Document 1 cannot be manufactured easily. For example, a gradient magnetic field coil is manufactured by separately manufacturing the center portion, end portion and step portion of the gradient magnetic field coil and positioning and connecting many wires formed in these parts to one another. Such a manufacturing process is complicated, and the manufacturing cost is increased.