Phase contrast magnetic resonance imaging

Atoms with an odd number of protons or neutrons have a randomly aligned angular spin momentum. When placed in a strong magnetic field, some of these spins align with the axis of the external field, which causes a net 'longitudinal' magnetization. These spins precess about the axis of the external field at a frequency proportional to the strength of that field. Then, energy is added to the system through a Radio frequency (RF) pulse to 'excite' the spins, changing the axis that the spins precess about. These spins can then be observed by receiver coils (Radiofrequency coils) using Faraday's law of induction. Different tissues respond to the added energy in different ways, and imaging parameters can be adjusted to highlight desired tissues.

All of these spins have a phase that is dependent on the atom's velocity. Phase shift ${\displaystyle (\phi )}$ of a spin is a function of the gradient field ${\displaystyle \mathbf {G} (t)}$:

${\displaystyle \phi =\gamma \int _{0}^{t}B_{0}+\mathbf {G} (\tau )\cdot \mathbf {r} (\tau )d\tau }$

where ${\displaystyle \gamma }$ is the Gyromagnetic ratio and ${\displaystyle \mathbf {r} }$ is defined as:

${\displaystyle \mathbf {r} (\tau )=\mathbf {r} _{0}+\mathbf {v} _{r}\tau +{\frac {1}{2}}\mathbf {a} _{r}\tau ^{2}+\ldots }$ ,

${\displaystyle \mathbf {r} _{0}}$ is the initial position of the spin, ${\displaystyle \mathbf {v} _{r}}$ is the spin velocity, and ${\displaystyle \mathbf {a} _{r}}$ is the spin acceleration.

If we only consider static spins and spins in the x-direction, we can rewrite equation for phase shift as:

${\displaystyle \phi =\gamma x_{0}\int _{0}^{t}G_{x}(\tau )d\tau +\gamma v_{x}\int _{0}^{t}G_{x}(\tau )\tau d\tau +\gamma {\frac {a_{x}}{2}}\int _{0}^{t}G_{x}(\tau )\tau ^{2}d\tau +\ldots }$

We then assume that acceleration and higher order terms are negligible to simplify the expression for phase to:

${\displaystyle \phi =\gamma (x_{0}M_{0}+v_{x}M_{1})}$

where ${\displaystyle M_{0}}$ is the zeroth moment of the x-gradient and ${\displaystyle M_{1}}$ is the first moment of the x gradient.

If we take two different acquisitions with applied magnetic gradients that are the opposite of each other (bipolar gradients), we can add the results of the two acquisitions together to calculate a change in phase that is dependent on gradient:

${\displaystyle \Delta \phi =v(\gamma \Delta M_{1})}$

where ${\displaystyle \Delta M_{1}=2M_{1}}$.[3] [4]

The phase shift is measured and converted to a velocity according to the following equation:

${\displaystyle v={\frac {v_{enc}}{\pi }}\Delta \phi }$

where ${\displaystyle v_{enc}}$ is the maximum velocity that can be recorded and ${\displaystyle \Delta \phi }$ is the recorded phase shift.

The choice of ${\displaystyle v_{enc}}$ defines range of velocities visible, known as the ‘dynamic range’. A choice of ${\displaystyle v_{enc}}$ below the maximum velocity in the slice will induce aliasing in the image where a velocity just greater than ${\displaystyle v_{enc}}$ will be incorrectly calculated as moving in the opposite direction. However, there is a direct trade-off between the maximum velocity that can be encoded and the signal-to-noise ratio of the velocity measurements. This can be described by:

${\displaystyle SNR_{v}={\frac {\pi }{\sqrt {2}}}{\frac {v}{v_{enc}}}SNR}$

where ${\displaystyle SNR}$ is the signal-to-noise ratio of the image (which depends on the magnetic field of the scanner, the voxel volume, and the acquisition time of the scan).

For an example, setting a ‘low’ ${\displaystyle v_{enc}}$ (below the maximum velocity expected in the scan) will allow for better visualization of slower velocities (better SNR), but any higher velocities will alias to an incorrect value. Setting a ‘high’ ${\displaystyle v_{enc}}$ (above the maximum velocity expected in the scan) will allow for the proper velocity quantification, but the larger dynamic range will obscure the smaller velocity features as well as decrease SNR. Therefore, the setting of ${\displaystyle v_{enc}}$ will be application dependent and care must be exercised in the selection. In order to further allow for proper velocity quantification, especially in clinical applications where the velocity dynamic range of flow is high (e.g. blood flow velocities in vessels across the thoracoabdominal cavity), a dual-echo PC-MRI (DEPC) method with dual velocity encoding in the same repetition time has been developed.[5] The DEPC method does not only allow for proper velocity quantification, but also reduces the total acquisition time (especially when applied to 4D flow imaging) compared to a single-echo single-${\displaystyle v_{enc}}$ PC-MRI acquisition carried out at two separate ${\displaystyle v_{enc}}$ values.

To allow for more flexibility in selecting ${\displaystyle v_{enc}}$, instantaneous phase (phase unwrapping) can be used to increase both dynamic range and SNR.[6]

Phase contrast MRI is one of the main techniques for magnetic resonance angiography (MRA). This is used to generate images of arteries (and less commonly veins) in order to evaluate them for stenosis (abnormal narrowing), occlusions, aneurysms (vessel wall dilatations, at risk of rupture) or other abnormalities. MRA is often used to evaluate the arteries of the neck and brain, the thoracic and abdominal aorta, the renal arteries, and the legs (the latter exam is often referred to as a "run-off").

In particular, a few limitations of PC-MRI are of importance for the measured velocities:

• Partial volume effects (when a voxel contains the boundary between static and moving materials) can overestimate phase leading to inaccurate velocities at the interface between materials or tissues.
• Intravoxel phase dispersion (when velocities within a pixel are heterogeneous or in areas of turbulent flow) can produce a resultant phase that does not resolve the flow features accurately.
• Assuming that acceleration and higher orders of motion are negligible can be inaccurate depending on the flow field.
• Displacement artifacts (also known as misregistration and oblique flow artifacts) occur when there is a time difference between the phase and frequency encoding. These artifacts are highest when the flow direction is within the slice plane (most prominent in the heart and aorta for biological flows)[10]

A Vastly undersampled Isotropic Projection Reconstruction (VIPR) is a radially acquired MRI sequence which results in high-resolution MRA with significantly reduced scan times, and without the need for breath-holding.[11]