Magnetic resonance imaging (MRI) is a powerful and versatile modality for preclinical studies. A particular strength of MRI is the wide variety of different image “contrasts”, many of which do not involve use of external contrast agents, that are available in imaging studies. These contrasts derive from the rich physics associated with the interaction of nuclear spins with external magnetic fields. This talk will provide an introduction to these physical principles.

The presentation will be didactic in nature, with an emphasis on principles and insights, rather than equations and mathematics. Many nuclei have the inherent physical property of “spin”, making them active for nuclear magnetic resonance (NMR) spectroscopy and imaging. Here, we focus on spin-1/2 nuclei, including ¹H, ³¹P, and ¹³C. In the absence of an external magnetic field, such spins-1/2 have two degenerate energy levels. In the presence of a field, these levels split into a low-energy state (spin aligned “parallel” to the field) and a high-energy state (spin aligned “anti-parallel” to the field) to the field. The lower energy state is preferentially populated, according to the Boltzmann factor. For an ensemble of nuclear spins, the difference between the number of spins parallel and anti-parallel to the external field generates a net magnetic moment (“magnetization”) that represents the available signal for MR spectroscopy and imaging experiments. As the strength of the external field increases, the energy difference between parallel and anti-parallel spin alignments increases. This drives an increase in the population difference between spin states, resulting in an increase in the magnitude of the net magnetic moment and, thus, signal intensity. This is one of the factors driving the push to the development of higher magnetic field scanners.

In general, detectors on MR scanners are unable to directly detect magnetization that is oriented along the direction of the applied magnetic field, designated by convention as the “z” direction. Instead, an MR experiment generally begins with the application of one or more radio-frequency (RF) pulses that nutate the net magnetization into the (“x-y”) plane perpendicular to the static field. The magnetization then precesses about the external field at a frequency (the so-called Larmor frequency) that is directly proportional to the field strength. The constant of proportionality, designated as the magnetogyric ratio, is nucleus specific.

As magnetization precesses in the x-y plane, variations in the precession frequency for different groups of spins will lead to a loss of phase coherence (“fanning out” of the magnetization vector), with a corresponding loss of MR signal. This “transverse relaxation” process is describable by time constants T2 or T2*. At the same time, the nuclear spins interact with their environment (the “lattice”), causing the net magnetization to return to its equilibrium direction along the static magnetic field. This “longitudinal relaxation” process is describable by the time constant T1. Together, these processes reflect a return to the state of thermal equilibrium and are globally referred to as relaxation phenomena: (i) the loss of signal due to loss of phase coherence amongst the vast number of spins that make up the macroscopic magnetic moment is referred to as transverse, or T2, or spin-spin relaxation; (ii) the restoration of the macroscopic magnetic moment along the direction of the static field is referred to as longitudinal, T1, or spin-lattice relaxation. Differences in T1 and T2 relaxation time constants amongst different tissues, and between healthy and pathologic tissue, are routinely used to generate image contrast in MRI experiments.

As noted above, the observed MR (Larmor) frequency at which nuclear spins precess about a magnetic field is directly proportional to the strength of the field. Thus, for nuclear spins (e.g., ¹H) in a single molecular species (e.g., water) in a uniform magnetic field, MR signal will be observed at a single, well-defined frequency. A single resonance line whose frequency is the same at all positions within a sample is incommensurate with MR as an imaging technique. Creating a dependence of MR signal frequency on position, as required for imaging, requires the use of magnetic-field gradients. Applying a strong, constant gradient in a specific, well-defined spatial direction creates a linear change in the magnetic-field strength (a linear function of spatial coordinate) alone the gradient axis. This results in the one-to-one correspondence between MR frequency and position necessary to produce a one-dimensional image. These considerations are easily generalized to the generation of three- dimensional images through the application of combinations of gradients in three orthogonal directions. MR data are typically collected in the presence of applied magnetic-field gradients, in what is known as spatial-frequency, or “k” space. In conventional 2D imaging, trajectories in k-space are generally represented by parallel lines in the kx, ky plane. Multidimensional Fourier transformation is then used to convert these k-space data into spatial images. Rapid acquisition of MR images (e.g., echo-planar, fast spin-echo, spiral) involve data collection employing alternate k-space trajectories.

Dephasing of magnetization in the transverse, x-y plane due to applied magnetic-field gradients or static inhomogeneities in the local magnetic field can be reversed, causing the magnetization to rephase and resulting in the formation of an “echo”. Echoes resulting from reversing the sign of the applied magnetic-field gradient are designated as “gradient echoes”, while those resulting from the application of one or more RF pulses are labeled “spin echoes”. Echoes provide a particularly efficient way of sampling k-space and are a standard component of most MR imaging sequences.

The vast majority of clinical and preclinical MR imaging studies involve the detection of water. (“It’s all about water!”) Water is ubiquitous in tissue – its high concentration provides the necessary sensitivity for the detection of MR signal – and its MR properties, including relaxation, are extremely sensitive to its local environment, contributing to image contrast. As noted above, differences in relaxation time constants T1 and/or T2 can serve as a powerful source of contrast amongst organs and tissues. Standard, anatomic imaging methods take advantage of these differences to generate “T1-weighted” and “T2-weighted” anatomic images that are the workhorses of MR imaging studies.

MR contrast agents are compounds, typically built around paramagnetic metal centers (e.g., Gd, Fe) that alter the relaxation properties of accessible water. Unlike the use of positron emission tomography (PET) tracers or optical agents, image contrast is generated not by the direct observation of the contrast agent, but, instead, through the effect of this agent on the relaxation properties of observed water.

In addition to relaxation parameters T1 and T2/T2*, image contrast can be made sensitive to many other properties of water, including:

Diffusion – Image contrast serves to measure the incoherent motion of water within tissue.

Perfusion/Flow – Image contrast reflects coherent flow.

Blood Oxygenation Level Dependence (BOLD) – Image contrast is sensitive to the balance between diamagnetic oxyhemoglobin and paramagnetic deoxyhemoglobin.

Magnetization Transfer (MT) – Image contrast reflects transfer of magnetization between water and exchangeable protons on neighboring macromolecules.