Magnetic Resonance Imaging

 

A detailed description can be found starting at this page. A synopsis follows.

 

Magnetic resonance imaging (MRI) can be distinguished in structural MRI which allows to obtain a structural image and functional MRI (fMRI) which when applied for instance to the brain, provides a functional profile representing brain activity over time. MRI uses nuclear magnetic resonance (NMR) and specifically it images hydrogen nuclei or protons which are very abundant in tissue (present in water molecules).

 

When a subject is placed in the strong magnetic field of an MRI scanner, usually 1.5-3 Tesla (the Earth’s magnetic field is 50μT), the magnetic field with exert a torque/force on the spins, causing their alignment with the direction of the magnetic field (conventionally on the z axis). The progressive alignment of an increasing number of spins results in the growth of a longitudinal magnetization along the z axis (Mz). This constitutes the first step of MRI, the process of polarization (spins become polarized) which is mediated in the clinical setting using a strong magnetic field.

 

If we apply a perpendicular radiofrequency (RF) pulse at the frequency of precession of the spins, termed Larmor frequency, the spins will absorb the RF energy and the magnetic field component of the RF will exert a torque/force on the longitudinal magnetization Mz leading to its tipping. Applying the RF pulse for a π/2 duration, will induce its tipping by 90° and its conversion to a transverse magnetization Mxy. For a 1.5 Tesla MRI system, the Larmor frequency for hydrogen nuclei is 63.85 MHz and therefore this radiofrequency is used.

 

The transverse magnetization Mxy will precess at the Larmor frequency upon the effect of the RF. Due to the prior spin polarization mediated by the magnetic field of the scanner, spins will be precessing in a synchronized manner demonstrating phase coherence. 

 

Following the end of the RF pulse, the energy absorption phase of the spins or the spin excitation phase is completed and an energy emission phase (relaxation) is triggered. Individual spins will be influenced by the magnetic effect of neighbouring spins, by potential collisions with them as well as by inhomogeneities of the magnetic field and they will start dephasing. In other words, as decelerations or phase “lags” may be induced, the phase of their rotational motion will be altered and practically they will not be “in phase” with other spins. As a result, a progressively decreasing number of spins will precess with the same phase, in a coherent manner and the Mxy magnetisation will decay exponentially in what is termed as T2 relaxation or spin-spin relaxation with T2 representing a time constant for this process. During this, the emitted energy which constitutes the signal of the MRI is captured by a coil. 

 

If we considered hypothetically that during the emission phase the Mxy magnetization was sustained, then we could imagine that each precession of its vector near a coil would generate a voltage in the form of a sine wave with frequency identical to the Larmor frequency and a stable amplitude. However, due to the decay of the Mxy magnetization which consists of the decrease of the number of spins precessing with the same phase, a dampening will be introduced and the generated signal (Figure 2) termed "Free Induction Decay" will be an exponentially damped sine wave. 

 

Different tissues have different T2. A long T2 provides a persistent Mxy magnetization and therefore a bright signal from Mxy. Conversely, a short T2 provides a faint signal. An MRI image will be a composite of different signal intensities representing different tissues. 

 

After completion of energy emission, the spins which are found in the xy plane, return (relax) to their original orientation in the z axis. This process is termed T1 relaxation or spin-lattice relaxation with T1 being a time constant for this process. Different tissues have different T1. A short T1 provides a quick recovery of the Mz magnetization to its original value and therefore a strong signal from Mz. A long T1 provides a long recovery of Mz magnetization and therefore a fainter signal from Mz. 

 

It is possible to create MRI images based on either T1 or T2 and these are called respectively T1-weighted or T2-weighted images. 

 

In conclusion, following the polarization phase, which constitutes the first step of MRI and which in the clinical setting is mediated using a strong magnetic field, the subsequent step is the detection phase. This phase includes, firstly, excitation with a radiofrequency to the purpose of signal generation and secondly, signal acquisition using a coil (it is noted that the subject continues to be in the same strong magnetic field).

 

MRI uses pulse sequences to mediate repeated spin excitation and relaxation accompanied by signal generation. Pulse sequences can mediate an off-on magnetization or a steady state magnetization which is termed steady state free precession (SSFP) and is linked to a steady state signal emission from the subject.

 

If the end of the 90° pulse and the decay of Mxy magnetization is followed by a 180° pulse, the spins will refocus again on the xy plane but in the opposite orientation and Mxy will regrow providing a new signal (or a signal repetition) termed spin echo (SE). If the echo is the result of three pulses e.g. 90°-90°-90° it is called a stimulated echo (STE). Additionally, it is possible to use magnetic field gradients in order to accelerate the FID and refocus the spins generating a gradient echo (GRE)

 

Steady state precession imaging is an MRI technique which uses steady state magnetization (SSFP). In general, SSFP sequences are based on gradient echo (GRE) sequences. They use dephasing and rephasing magnetic field gradients to refocus and record the FID (SSFP-FID sequences), the spin echo/stimulated spin echo (SSFP-SE/STE) or both (SSFP-Double and SSFP-Balanced sequences).