User:GyroMagician/MRI hardware

Main Magnet

FID signal from a badly shimmed sample has a complex envelope.

FID signal from a well shimmed sample, showing a pure exponential decay.

Shims

When a sample is placed into the scanner, the main magnetic field is distorted by susceptibility boundaries within that sample, causing signal dropout (regions showing no signal) and spatial distortions in acquired images. For humans or animals the effect is particularly pronounced at air-tissue boundaries such as the sinuses making, for example, the frontal lobes of the brain difficult image. To restore field homogeneity a set of shim coils are included in the scanner. These are resistive coils, usually at room temperature, capable of producing field corrections distributed as several orders of spherical harmonics.

After placing the sample in the scanner, the B0 field is ‘shimmed’ by adjusting currents in the shim coils. Field homogeneity is measured by examining an FID signal in the absence of field gradients. The FID from a poorly shimmed sample will show a complex decay envelope, often with many humps. Shim currents are then adjusted to produce a large amplitude exponentially decaying FID, indicating a homogeneous B0 field. The process is usually automated.

Gradients

Gradient coils are used to spatially encode the source of the NMR signal. Applying a magnetic field gradient across the imaging volume cause the Larmor frequency to vary as a function of position across the sample. A typical imaging system has three gradient coils, designed to produce orthogonal, highly linear field gradients across the imaging volume. The magnetic field generated by each coil is aligned to the main static field, but varies in the three orthogonal directions x, y and z. By combining field gradients, images can be acquired in any plane.

Gradient coils are large resistive electromagnets, set inside the bore of the main magnet. Due to the high current carried by the coils during an imaging sequence (~600A), they are often water cooled to prevent over-heating. Typical field gradients are in the range 20 to 100mT/m (i.e. in a 1.5 T magnet, when a maximal z-axis gradient is applied the field strength may be 1.45 T at one end of a 1 m long bore and 1.55 T at the other).

The slew rate of a gradient system is a measure of how quickly the gradients can be ramped on or off. Typical high performance gradients have a slew rate of up to 100-200 T/m/s. Slew rate depends both on the inductance of the gradient coil and on the power of the gradient amplifier.

Scan speed is dependent on performance of the gradient system. Stronger gradients allow for faster imaging, or for higher spatial resolution; similarly, gradients systems capable of faster switching can also permit faster scanning. In practice, the speed and strength of the gradient system are limited by the patient, rather than technology. Faster, stronger gradients produce more acoustic noise, and generate stronger electric fields within the body, leading to peripheral nerve stimulation.

A gradient coil creates an additional, linearly varying magnetic field that adds or subtracts from the main magnetic field. This additional magnetic field will have components in all 3 axes; however, only the component along the magnetic field (usually called the Z-axis, hence denoted $G_{z}$) is useful for imaging. Along any given axis, the gradient will add to the magnetic field on one side of the zero position and subtract from it on the other side.

Since the additional field is a gradient, it has units of millitesla per meter (mT/m).

During an imaging sequence the static B₀ field is deliberately altered by the addition of linear field gradients used to localize the NMR signal. An MRI system typically has three orthogonal gradient coils.

Early imaging experiments used shim coils used to generate field gradients [13], but it was quickly found that the imaging speed was limited by the time taken to switch the shim coils. Separate, low-inductance higher-power gradient coils are now used.