Current carrying magnet coils are used for a variety of purposes in NMR imaging apparatus. Examples include large electro-magnets designed to provide static magnetic fields to polarize nuclear spins, magnetic field gradient coils which superimpose gradients onto the static polarizing field and rf transmitter and receiver coils.
In many cases the design of a magnet coil is such as to optimize the magnetic field within a desired volume. However the coil inevitably produces an extraneous magnetic field outside that volume, especially relatively close to the coil. In the case of large bore static electromagnets the high fields they generate will produce undesirably strong extraneous fields at distances outside the magnet corresponding to many magnet diameters. Such magnet systems therefore require much free and unusable space around their installation. Stray magnetic fields may be intolerable in hospitals because of iron structures positioned around the installation site which vitiate the homogeneity of the magnetic field. Additionally, electronic equipment may not perform well in an environment which has an extraneous magnetic field.
Furthermore, most NMR imaging systems utilize rapidly switched magnetic field gradients in their operation. A major problem epecially where super-conductive magnets are used, is the interaction of the gradient field with the magnet itself. Existing attempts to minimize this interaction include the use of conducting metal screening sleeves. However, induced currents in these sleeves or in the heat shield of the magnet decay with uncontrolled relaxation times which make it difficult or even impossible to implement some of the faster and more efficient NMR imaging methods. This is because the decaying current produces image fields superimposed on the desired gradient field. This uncontrolled time dependence introduces phase artefacts which can completely ruin the image.
In order to provide adequate access for patients, and to improve gradient uniformity, it is desirable to maximize the diameter of the magnetic field gradient coils in an NMR imaging machine. However, this often causes the coils to be close to other conductors, either the surfaces of cryogenic vessels (in superconducting magnet systems), electromagnetic coil supports (in resistive magnet systems), or ferromagnetic pole pieces (in ferromagnetic systems). When gradients are switched rapidly, as with many imaging techniques, eddy currents are induced in these conductors which then contribute additional field gradients varying in time and potentially very non-uniform in space. Typical time constants for the decay of the eddy currents vary from a few milliseconds to hundreds of milliseconds, depending on the type of main magnet and the specific coil configuration.
The commonest solution to this problem is to tailor the input applied to the amplifiers generating the gradient coil currents in such a way that the gradient fields themselves follow the prescribed time variation. The input voltage and gradient coil currents are characteristically over-driven for the initial part of the on-period of the gradient. But this remedy has a major disadvantage. If the gradient coils are placed close to the coupled surfaces, so that the eddy current field gradients may have the same uniformity as the desired gradient, the gradient coils become very inefficient and a large over capacity in the gradient current amplifiers is required, since the `reflected` fields will be large and in the opposite sense from the desired fields. If, on the other hand, the gradient coils are reduced in size, to avoid the amplifier capacity problem, then the reflected gradient fields will in general be non-linear over the region of interest. Furthermore, in either case there are likely to be reflected fields from more distant conductors in the main magnet structure, each with its distinct time constant and spatial variation.
The only effective solution is in some way to reduce the gradient fields to zero at a finite radius outside be coupled to them.
Partially effective methods for magnetic screening in specific coil geometries have been proposed hitherto in particular U.S. Pat. Nos. 3,466,499 and 3,671,902. These geometries are not generally useful in NMR and NMR imaging.
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