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.

(SE) The most common pulse sequence used in MR imaging is based of the detection of a spin or Hahn echo. It uses 90° radio frequency pulses to excite the magnetization and one or more 180° pulses to refocus the spins to generate signal echoes named spin echoes (SE).
In the pulse sequence timing diagram, the simplest form of a spin echo sequence is illustrated.
The 90° excitation pulse rotates the longitudinal magnetization (Mz) into the xy-plane and the dephasing of the transverse magnetization (Mxy) starts.
The following application of a 180° refocusing pulse (rotates the magnetization in the x-plane) generates signal echoes. The purpose of the 180° pulse is to rephase the spins, causing them to regain coherence and thereby to recover transverse magnetization, producing a spin echo.
The recovery of the z-magnetization occurs with the T1 relaxation time and typically at a much slower rate than the T2-decay, because in general T1 is greater than T2 for living tissues and is in the range of 100-2000 ms.
The SE pulse sequence was devised in the early days of NMR days by Carr and Purcell and exists now in many forms: the multi echo pulse sequence using single or multislice acquisition, the fast spin echo (FSE/TSE) pulse sequence, echo planar imaging (EPI) pulse sequence and the gradient and spin echo (GRASE) pulse sequence;; all are basically spin echo sequences.
In the simplest form of SE imaging, the pulse sequence has to be repeated as many times as the image has lines.
Contrast values:
PD weighted: Short TE (20 ms) and long TR.
T1 weighted: Short TE (10-20 ms) and short TR (300-600 ms)
T2 weighted: Long TE (greater than 60 ms) and long TR (greater than 1600 ms)
With spin echo imaging no T2* occurs, caused by the 180° refocusing pulse. For this reason, spin echo sequences are more robust against e.g., susceptibility artifacts than gradient echo sequences.

See also Pulse Sequence Timing Diagram to find a description of the components.

(SFP or SSFP) Steady state free precession is any field or gradient echo sequence in which a non-zero steady state develops for both components of magnetization (transverse and longitudinal) and also a condition where the TR is shorter than the T1 and T2 times of the tissue. If the RF pulses are close enough together, the MR signal will never completely decay, implying that the spins in the transverse plane never completely dephase. The flip angle and the TR maintain the steady state. The flip angle should be 60-90° if the TR is 100 ms, if the TR is less than 100 ms, then the flip angle for steady state should be 45-60°.
Steady state free precession is also a method of MR excitation in which strings of RF pulses are applied rapidly and repeatedly with interpulse intervals short compared to both T1 and T2. Alternating the phases of the RF pulses by 180° can be useful. The signal reforms as an echo immediately before each RF pulse; immediately after the RF pulse there is additional signal from the FID produced by the pulse.
The strength of the FID will depend on the time between pulses (TR), the tissue and the flip angle of the pulse; the strength of the echo will additionally depend on the T2 of the tissue. With the use of appropriate dephasing gradients, the signal can be observed as a frequency-encoded gradient echo either shortly before the RF pulse or after it; the signal immediately before the RF pulse will be more highly T2 weighted. The signal immediately after the RF pulse (in a rapid series of RF pulses) will depend on T2 as well as T1, unless measures are taken to destroy signal refocusing and prevent the development of steady state free precession.
To avoid setting up a state of SSFP when using rapidly repeated excitation RF pulses, it may be necessary to spoil the phase coherence between excitations, e.g. with varying phase shifts or timing of the exciting RF pulses or varying spoiler gradient pulses between the excitations.
Steady state free precession imaging methods are quite sensitive to the resonant frequency of the material. Fluctuating equilibrium MR (see also FIESTA and DRIVE)and linear combination SSFP actually use this sensitivity for fat suppression. Fat saturated SSFP (FS-SSFP) use a more complex fat suppression scheme than FEMR or LCSSFP, but has a 40% lower scan time.
A new family of steady state free precession sequences use a balanced gradient, a gradient waveform, which will act on any stationary spin on resonance between 2 consecutive RF pulses and return it to the same phase it had before the gradients were applied.
This sequences include, e.g. Balanced Fast Field Echo - bFFE, Balanced Turbo Field Echo - bTFE, Fast Imaging with Steady Precession - TrueFISP and Balanced SARGE - BASG.

See also FIESTA.

The dephasing of the protons is named the T2, spin-spin or transverse relaxation. The T2 time constant is the time taken for spinning protons to lose phase coherence among the nuclei spinning perpendicular to the main field. This interaction between spins results in a reduction in the transverse magnetization. The value of T2 depends on the mobility of the protons. A large mobility results in an average magnetic field variation of zero, resulting in a long T2 period of this tissue.

See also T2 Time.

(T2* or T two star) The observed time constant of the FID due to loss of phase coherence among spins oriented at an angle to the static magnetic field. Commonly due to a combination of magnetic field inhomogeneities, dB, and spin spin transverse relaxation, with the result of rapid loss in transverse magnetization and MRI signal. MRI signals can usually still be recovered as a spin echo in times less than or on the order of T2.
1/T2 * @ 1/T2 + Dw/2; Dw = gDB. The FID will generally not be exponential, so that T2* will not be unique.