Echo planar imaging (EPI), proposed by Mansfield in 1977, is a much faster imaging technique, which aims at generating a series of spin or gradient echoes with different phase encoding in the short period of time after a single excitation of the spin system by an RF pulse.
In the EPI sequence, image acquisition typically begins with a standard 900 flip, negative lobes of the phase encoding and frequency encoding gradients then cause a phase evolution that moves the spin system to the periphery of the k-space. The frequency encoding gradient or read out gradient is then rapidly switched between maximum positive and negative amplitude, repeatedly refocusing the transverse magnetization to yield a train of gradient echoes while sweeping across kxin alternate
directions. Alternating k-space lines (echoes) are traversed in opposite directions of the readout gradient, therefore every other line has to be time-reversed before Fourier transform is applied to generate an image. The effective echo time TE occurs when the maximum amplitude of the echoes occurs. In the EPI technique acquisition of the data must proceed in a period less than T2* (around 25 to 50ms), a schematic EPI
sequence based on gradient echo shown in Figure 1.28. EPI is good for a quantification study is the and most commonly used dynamic sequence.
Figure 1.28 Schematic diagram of the echo planar imaging and k-space trajectory. The repeatedly reversing readout gradients and blipped phase encoding gradients forms a series of echoes.
The sensitivity of EPI sequence images can be increased with high field strength scanners, however, the implementation of EPI in the high magnetic field gives rise to a number of challenges. In EPI, the acquisition time for a given spatial resolution is proportional to the field of view in the phase encoding direction and thus to the number of k-space lines. Shortening of T2 and T2* poses a limit on the number of k- space lines that can be acquired after excitation, as it is not possible to sample the entire k-space with a long echo train length and retain a good TE. Higher resolution also puts demands on the gradients and due to the use of high slew rates can cause peripheral nerve stimulation (PNS) in the subject.
Inhomogeneities of the B0field due to poor shimming and susceptibility effects influence
the evolution of the phase during the echo train, affecting EPI severely. Local field inhomogeneities and chemical shifts, results geometric distortions and phase variations that accumulates along the phase encode direction and can cause substantial displacement of the signal during EPI. Imperfections in the gradient coils or eddy currents cause a slight offset between lines collected in the positive and negative directions, leading to a Nyquist ghosting artefact in the reconstructed image.
1.2.11 Artefacts
In MRI, artefacts manifest as positive or negative signal intensities that do not accurately represent the imaged subject. Although some artefacts are relatively
insignificant and easily identified, others can limit the diagnostic potential of the exam by obscuring the images. An overview of MRI artefacts is discussed below. Motion artefacts:
The most ubiquitous and noticeable artefacts in MRI arise with patient motion, that includes voluntary and involuntary movement and flow (blood, CSF). The long acquisition time of certain MRI sequences increases the probability of motion blurring and contrast resolution losses. Motion artefacts mostly occur along the phase encode direction, since the phase encoding lines are separated by a TR interval that can last for 3 seconds or longer. Thus even very slight motion will cause a change in the recorded phase variation across the FOV throughout the acquisition. The frequency encode direction is less affected especially by periodic motion, since the evolution of the echo signal, frequency encoding, and sampling occur simultaneously over several milliseconds. There are several techniques for motion artefact corrections, such as,
-cardiac and respiratory gating
-signal averaging to reduce artefacts of random motion -short TE that limits phase evaluation during the readout
-gradient moment nulling by additional gradient pulses for flow compensation -presaturation pulses applied outside the imaging region to suppress flowing blood
Machine dependent artefacts:
Magnetic field inhomogeneities are either global or local field perturbations that lead to the mismapping of tissues within the image, and can cause more rapid T2* relaxation. Automatic shimming helps to reduce the magnetic field inhomogeneities. Water fat shift:
The water-fat shift arises due to the chemical shift of fat with respect to water occurs in measurement direction, and it depends on the frequency bandwidth of the measurement gradient. The higher the bandwidth the smaller the WFS and lowers the SNR. WFS is field strength dependent, larger field strength will result in larger WFS. Fat suppression: STIR
STIR: TheShort tau inversion recovery (STIR) pulse sequence uses a very short TI for magnitude signal acquisition. As the inversion recovery (IR) sequence based on recovery from the transverse magnetization, it is possible to separate specific tissues (such as fat) by varying inversion time (TI). When the signal from the target tissue
passes the null point the 900 pulse is applied to create the echo of the whole spin system, while no signal from the target tissue.
Partial volume artefacts:
The finite resolution of the imaging device generally results in a single voxel representing more than one tissue type, this is known as the partial volume effect (PVE). Partial volume artefacts arise from the finite size of the voxel over which the signal is averaged. This results in a loss of detail and spatial resolution. Reduction of partial volume artefacts can be undertaken by using a smaller pixel size and/or a smaller slice thickness.
1.2.12 Instrumentation
The MRI scanner is composed of multitude of components, the main magnet, the gradient coils, the RF system, the console controlling the scanner and different electronics parts. This section provides a brief discussion of the MRI instruments. Main magnet: The central component of an MR scanner is the main magnet which produces the B0 field. The high magnetic field is usually produced by an
electromagnet made from coils of niobium-titanium (Nb3Ti) wire, which become
superconducting at about 10K (-2630C). To produce such a low temperature, a bath of liquid helium is used to keep the wires superconducting. Active or passive shielding also surrounds the magnet in order to reduce the fringe of the magnetic field at the edge of the magnet.
Shim coils: Homogeneity of the B0 field is very important in minimizing the spin
dephasing. Susceptibility differences within the sample itself also introduce spatial variations, so that the field needs to be shimmed dynamically at the beginning of each scan session. The MR scanner incorporates shim coils that produce compensatory magnetic fields to correct for spatial variations in the main magnetic field. This is done to ensure that the generated magnetic field is homogeneous enough for imaging purposes.
Gradient coils: An MR scanner contains three orthogonal gradient coils, which produce magnetic fields that vary linearly in strength along each direction. The gradient coils are generally placed inside the shim coils and deliberately alter the magnetic field when energized in order to localize the MR signal. They can be used in combination to produce magnetic field gradients in any direction, and allow images to be acquired in arbitrary planes. For this thesis data acquired from the Philips 3.0T
Achieva scanner, Philips Medical Systems, Best, the Netherlands; the scanner is set up with three fully independent gradient axes for orthogonal, oblique and double- oblique imaging with maximum amplitude 40 mT/m, maximum slew rate 120 mT/m/ms, minimum imaging rise time 0.33 ms, these are non-resonant, actively shielded and 100% duty cycle gradients.
RF coils: The MR scanner includes RF transmitter and related circuitry to produce the