A coil is a large inductor with a considerable dimension and a defined wavelength, commonly used in configurations for MR imaging. The frequency of the radio frequency coil is defined by the Larmor relationship.
The MRI image quality depends on the signal to noise ratio (SNR) of the acquired signal from the patient. Several MR imaging coils are necessary to handle the diversity of applications. Large coils have a large measurement field, but low signal intensity and vice versa (see also coil diameter). The closer the coil to the object, the stronger the signal - the smaller the volume, the higher the SNR. SNR is very important in obtaining clear images of the human body. The shape of the coil depends on the image sampling. The best available homogeneity can be reached by choice of the appropriate coil type and correct coil positioning. Orientation is critical to the sensitivity of the RF coil and therefore the coil should be perpendicular to the static magnetic field.

RF coils can be differentiated by there function into three general categories:
The RF signal is in the range of 10 to 100 MHz. During a typical set of clinical image measurements, the entire frequency spectrum of interest is of the order 10 kHz, which is an extremely narrow band, considering that the center frequency is about 100 MHz. This allows the use of single-frequency matching techniques for coils because their inherent bandwidth always exceeds the image bandwidth. The multi turn solenoid, bird cage coil, single turn solenoid, and saddle coil are typically operated as the transmitter and receiver of RF energy. The surface and phased array coils are typically operated as a receive only coil.

See also the related poll result: '3rd party coils are better than the original manufacturer coils'

A surface coil placed over a region of interest will have an effective selectivity for a volume approximately subtended by the coil circumference and one radius deep from the coil center. Such a coil can be used for simple localization of sites for measurement of chemical shift spectra, especially of phosphorus, and blood flow studies. Some additional spatial selectivity can be achieved with magnetic field gradients.

MRI coils with a small diameter obtain a higher signal to noise ratio (SNR) than coils with a large diameter. A surface coil with a small diameter can be used to improve the resolution because the area of interest is around the optimal signal depth. The field of view of a (superficial) surface coil is half the diameter of the coil. A disadvantage is a lower sensitive volume of the coil. By combining several coils with small diameters (phased array coil) to record the signal simultaneously and independently, the SNR level improves considerably.

Process by which regions of tissue are selectively sampled to produce spectra from defined volumes in space. These methods may be employed to sample a single region in space (single voxel method) or multiple regions simultaneously (multivoxel methods). The spatial selectivity can be achieved by a variety of methods including surface coils, surface coils in conjunction with RF gradient methods, or RF pulses in combination with switched magnetic field gradients, for example, volume-selective excitation. An indirect method of achieving spatial selectivity is the destruction of coherence of the magnetization in regions that lie outside the region of interest. A variety of spatial encoding schemes have been employed for multivoxel localization. See Chemical shift imaging.

Quantification relies on inflow effects or on spin phase effects and therefore on quantifying the phase shifts of moving tissues relative to stationary tissues.
With properly designed pulse sequences (see phase contrast sequence) the pixel by pixel phase represents a map of the velocities measured in the imaging plane. Spin phase effect-based flow quantification schemes use pulse sequences specifically designed so that the phase angle in a pixel obtained upon measuring the signal is proportional to the velocity. As the relation of the phase angle to the velocity is defined by the gradient amplitudes and the gradient switch-on times, which are known, velocity can be determined quantitatively on a pixel-by-pixel basis. Once, this velocity is known, the flow in a vessel can be determined by multiplying the pixel area with the pixel velocity. Summing this quantity for all pixels inside a vessel results in a flow volume, which is measured, e.g. in ml/sec.
Flow related enhancement-based flow quantification techniques (entry phenomena) work because spins in a section perpendicular to the vessel of interest are labeled with some radio frequency RF pulse. Positional readout of the tagged spins some time T later will show the distance D they have traveled.
For constant flow, the velocity v is obtained by dividing the distance D by the time T : v = D/T. Variations of this basic principle have been proposed to measure flow, but the standard methods to measure velocity and flow use the spin phase effect.
Cardiac MRI sequences are used to encode images with velocity information. These pulse sequences permit quantification of flow-related physiologic data, such as blood flow in the aorta or pulmonary arteries and the peak velocity across stenotic valves.