Introduction

Magnetization is the cornerstone of magnetic resonance imaging (MRI) signal generation, but its role extends far beyond simple relaxation. In T2*-weighted imaging, the decay of the transverse magnetization is exquisitely sensitive to local magnetic field inhomogeneities caused by variations in tissue susceptibility. Unlike conventional T2-weighted sequences, which refocus static dephasing using a 180° refocusing pulse, T2* sequences rely on gradient echoes that allow those field perturbations to accumulate. This makes T2*-weighted imaging a powerful tool for detecting iron deposition, hemorrhage, calcification, and other susceptibility-related tissue changes. Understanding the underlying magnetization physics is essential for both radiologists and MR physicists to correctly interpret images and optimize acquisition parameters. This article provides a comprehensive, clinically oriented review of how magnetization governs susceptibility effects in T2*-weighted imaging.

Basics of Magnetization in MRI

Nuclear Spin and Net Magnetization

Hydrogen nuclei (1H) possess a magnetic moment due to their spin. When placed in a strong external magnetic field (B0), these moments align either parallel (lower energy) or antiparallel (higher energy) to the field. A slight excess in the parallel orientation creates a net magnetization vector (M) pointing along the longitudinal axis. The magnitude of this net magnetization depends on field strength, temperature, and the number of hydrogen spins per unit volume (proton density).

Excitation and Relaxation

Applying a resonant radiofrequency (RF) pulse tips M away from the longitudinal axis into the transverse plane. After the pulse, two independent relaxation processes occur:

  • T1 relaxation (spin–lattice): recovery of longitudinal magnetization via energy transfer to the surrounding lattice; governs how quickly the signal can be repeated.
  • T2 relaxation (spin–spin): decay of transverse magnetization due to random interactions between neighboring spins, causing irreversible dephasing.

In practice, additional dephasing occurs from static magnetic field inhomogeneities – both macroscopic (imperfect shim, air–tissue boundaries) and microscopic (varying susceptibilities within the tissue). The combined effect is described by a shorter time constant, T2* (read as T2-star). The relationship is given by: 1/T2* = 1/T2 + 1/T2′, where T2′ represents the component due to static field inhomogeneities.

Gradient Echo vs. Spin Echo

Spin echo (SE) sequences deploy a 180° refocusing pulse that reverses the phase accumulation caused by static field inhomogeneities, producing a signal envelope that decays with T2, not T2*. Gradient echo (GRE) sequences, on the other hand, omit the 180° pulse. Instead, they use a bipolar readout gradient to form an echo; any off-resonance effects accumulate over successive echoes, yielding a signal that decays with T2*. Therefore, GRE (and its derivative, susceptibility-weighted imaging) is the method of choice for probing local susceptibility variations.

T2*-Weighted Imaging and Susceptibility

What Is Magnetic Susceptibility?

Magnetic susceptibility (χ) is a dimensionless property describing how a material becomes magnetized when placed in a magnetic field. Materials can be classified as:

  • Diamagnetic (χ < 0): weakly oppose the field (e.g., water, calcium, bone).
  • Paramagnetic (χ > 0): weakly enhance the field (e.g., gadolinium, deoxyhemoglobin, iron in ferritin).
  • Ferromagnetic (χ ≫ 0): strongly enhance and remain magnetized (rare in vivo).
  • Superparamagnetic (χ intermediate): particles that become magnetized in a field but do not retain magnetization after removal (e.g., iron oxide nanoparticles used as contrast agents).

Differences in χ between adjacent tissues create local magnetic field gradients. In T2*-weighted GRE sequences, spins in these regions precess at different frequencies, causing rapid dephasing and signal loss – the hallmark of susceptibility effects.

T2* Contrast Mechanisms

The signal intensity in a T2*-weighted image reflects the local transverse relaxation rate. Strong paramagnetic substances (like hemosiderin from chronic hemorrhage) produce large field perturbations, resulting in profound signal loss. This is exploited in clinical imaging to identify microbleeds, iron overload, and venous structures (where deoxygenated blood is paramagnetic). Conversely, diamagnetic materials such as calcifications cause less signal drop but may appear hypointense relative to background tissue due to their different susceptibility.

Pure T2* weighting is achieved by selecting a long echo time (TE) in a GRE sequence, typically 15–40 ms at 1.5T and 10–30 ms at 3T, depending on the targeted tissue. At higher field strengths, the susceptibility effects are magnified proportionally to B0, offering improved contrast but also increased artifacts. For a detailed description of GRE sequence parameters, see the Radiopaedia article on gradient echo sequences.

Susceptibility-Weighted Imaging (SWI)

A specialized extension of T2*-weighted GRE is susceptibility-weighted imaging (SWI), which combines magnitude and phase information to enhance contrast for substances with different susceptibilities. SWI is highly sensitive for venous structures, microbleeds, and iron content, and has become a standard sequence in brain MRI protocols. The phase image provides additional information: paramagnetic substances cause a negative phase shift (relative to background) while diamagnetic substances cause a positive shift, helping to discriminate, for example, hemorrhage from calcification.

Magnetization and Susceptibility Effects

Local Field Distortions and Dephasing

When a tissue region contains a substance with susceptibility different from its surroundings, the magnetic field is perturbed both inside and outside the region. This perturbation, described by Maxwell's equations, creates field gradients that vary spatially. Spins diffusing through these gradients experience additional random dephasing – called dynamic dephasing – while stationary spins experience static dephasing that is partially reversible with a 180° pulse. In GRE sequences, both static and dynamic dephasing contribute to T2* decay.

The magnitude of the field offset at a point r is given (to first order) by:

ΔB(r) = (χ(r) – χref) · B0 · f(geometry)

where χref is the susceptibility of the reference medium (often water or brain parenchyma), and f(geometry) accounts for the shape of the susceptibility source (sphere, cylinder, etc.). For spherical iron deposits, the field perturbation decays as 1/r³ from the center, creating a characteristic dipole pattern.

Key Sources of Susceptibility Variation In Vivo

  • Iron: Stored as ferritin and hemosiderin in the liver, spleen, and deep brain nuclei (globus pallidus, substantia nigra, red nucleus, dentate nucleus). Excessive iron accumulation is seen in neurodegenerative diseases (Parkinson's, Alzheimer's, multiple sclerosis) and in hemochromatosis. The paramagnetic effect of iron shortens T2* and produces hypointensity on GRE images.
  • Deoxygenated blood: Deoxyhemoglobin is paramagnetic due to the four unpaired electrons in its heme iron. In venous blood, the paramagnetism creates phase differences between the vessel lumen and parenchyma, which is the basis for BOLD (blood oxygen level–dependent) imaging and for visualizing cerebral veins in SWI.
  • Calcium: Calcium is diamagnetic. In a GRE sequence it causes a small signal loss, but the phase shift is opposite to that of paramagnetic substances (i.e., positive phase shift relative to background). This polarity difference is exploited in SWI phase images to differentiate calcified and hemorrhagic lesions.
  • Air–tissue interfaces: The large susceptibility difference between air (χ ≈ 0.36×10⁻⁶, essentially zero) and tissue (χ ≈ –9×10⁻⁶) creates severe field gradients at sinuses, skull base, and lung–chest wall interfaces. This leads to signal dropout and geometric distortion, particularly in echo-planar imaging (EPI).
  • Contrast agents: Gadolinium-based agents are paramagnetic and shorten T1 (used for T1-weighted enhancement) but also affect T2* by causing local field inhomogeneities at high concentrations or in calcium-complexed forms (e.g., macrocyclic vs linear agents).

Susceptibility Artifacts in Clinical Imaging

While susceptibility effects are valuable for tissue characterization, they also cause image artifacts that degrade diagnostic quality. The most common are:

  • Signal dropout: Complete loss of signal in regions of rapid dephasing, e.g., adjacent to surgical clips, dental hardware, or the petrous bone.
  • Geometric distortion: Misregistration of spins due to frequency shifts, prominent in EPI sequences (diffusion-weighted imaging).
  • Blurring and ringing: Occurs when the point-spread function is broadened by field inhomogeneities, especially with long echo trains.

Strategies to mitigate these artifacts include improved shimming, shorter TE, smaller voxel size, parallel imaging, and the use of view-angle tilting or z-shimming. For a comprehensive review of artifact reduction techniques, see the article on susceptibility artifacts in MRI by Hargreaves et al. (Radiographics).

Quantitative Susceptibility Mapping (QSM)

A more recent development is quantitative susceptibility mapping (QSM), which reconstructs the underlying spatial distribution of magnetic susceptibility from the phase data of a GRE acquisition. QSM removes the nonlocal dipole effect and provides a direct measure of tissue iron content, calcium content, and other paramagnetic/diamagnetic species. It has found clinical applications in quantifying iron accumulation in neurodegenerative diseases, estimating hemorrhage volume in cerebral microbleeds, and in venography. For an in-depth explanation of QSM principles and clinical examples, the review by Wang and Liu (2015) in NMR in Biomedicine offers a thorough introduction.

Implications for Imaging and Diagnosis

Hemorrhage Detection (Cerebral Microbleeds)

T2*-weighted GRE and SWI are the sequences of choice for detecting cerebral microbleeds (CMBs). These small, chronic hemorrhages present as hypointense foci due to hemosiderin deposition. CMBs are associated with hypertension, cerebral amyloid angiopathy, traumatic brain injury, and anticoagulant therapy. Their presence and distribution (lobar vs deep) help differentiate the underlying small vessel disease. A typical brain MRI protocol for trauma or cognitive decline includes a GRE sequence with TE > 20 ms at 3T. The advent of SWI has increased sensitivity for CMBs by up to 6 times compared to conventional GRE.

Iron Overload Disorders

Patients with hereditary hemochromatosis, thalassemia, or repeated blood transfusions accumulate iron in the liver, heart, and endocrine organs. The T2* relaxation time in the liver correlates inversely with iron concentration. Liver T2* mapping using a multi-echo GRE sequence allows noninvasive quantification of hepatic iron content, helping guide chelation therapy. Similarly, myocardial T2* mapping (e.g., for patients with thalassemia) predicts the risk of cardiac failure – a T2* < 10 ms at 1.5T indicates severe iron overload and increased risk of arrhythmia.

Neurodegenerative Diseases

Abnormal iron deposition in deep gray matter nuclei is a hallmark of several neurodegenerative conditions. In Parkinson's disease, increased iron in the substantia nigra pars compacta correlates with motor symptom severity. In multiple sclerosis, iron rim lesions (paramagnetic rim lesions on SWI) indicate chronic, active plaques with ongoing inflammation. Alzheimer's disease shows elevated iron in the hippocampus and basal ganglia. QSM can provide quantitative iron maps that track disease progression and may serve as an imaging biomarker.

Calcification vs. Hemorrhage Differentiation

Both calcification and chronic hemorrhage appear hypointense on magnitude T2*-weighted images. However, phase images from SWI or QSM reveal opposite phase shifts: calcifications are diamagnetic (positive phase shift, i.e., higher frequency) while hemorrhage is paramagnetic (negative phase shift). This distinction is critical for characterizing brain tumors, phakomatoses (e.g., tuberous sclerosis), and vascular malformations. For example, differentiating a calcified cavernoma from a hemorrhagic metastasis guides further management.

Vascular Imaging and BOLD

T2*-weighted imaging is the basis of functional MRI (fMRI) using the BOLD effect. Deoxyhemoglobin in capillary beds acts as an endogenous contrast agent. During neuronal activation, blood flow increases, diluting deoxyhemoglobin, which reduces the paramagnetic dephasing and increases T2*-weighted signal. While BOLD fMRI is used predominantly for mapping brain function, the underlying susceptibility physics is identical to that of T2* contrast in structural imaging.

Beyond the brain, T2*-weighted sequences are employed in the heart for detecting myocardial iron, in the liver for fatty liver disease assessment (via T2* correction for IDEAL water–fat separation), and in musculoskeletal imaging for detecting hemosiderin in hemophilic arthropathy or pigmented villonodular synovitis (PVNS).

Practical Considerations for Technologists and Radiologists

Sequence Optimization

To obtain diagnostic T2*-weighted images, the following parameters should be tailored to the clinical question and field strength:

  • Echo time (TE): For SWI, use long TE to maximize susceptibility contrast but avoid excessive signal loss (typical TE at 1.5T: 40 ms; at 3T: 20–25 ms; at 7T: 10–15 ms).
  • Resolution: Use isotropic voxels (e.g., 0.5–1 mm³) for SWI to resolve small vessels and microbleeds; larger voxels increase partial volume effects and reduce contrast.
  • Flip angle: A small flip angle (15–25°) reduces T1 weighting and maximizes T2* contrast; a larger flip angle (≥30°) adds T1 weighting, which may be desirable for specific applications.
  • Flow compensation: Apply gradient moment nulling in the slice and readout directions to suppress pulsatile flow artifacts from arteries and veins.
  • Parallel imaging: Use acceleration factors of 2–3 to shorten scan time and reduce susceptibility-related blurring; higher acceleration may cause g-factor noise.
  • Shimming: Active shimming (particularly second-order) is critical at 3T and above to minimize macroscopic B0 inhomogeneities, especially near air–tissue interfaces. In regions like the orbit or temporal lobes, local shim may be required.

Artifact Mitigation in Clinical Practice

When susceptibility artifacts degrade image quality, consider the following adjustments:

  • Reduce TE to decrease dephasing time (at the cost of T2* contrast).
  • Increase receiver bandwidth and/or shorten echo train length for EPI-based sequences (e.g., DWI).
  • Use a single-shot or multi-shot GRE with segmented k-space readouts to reduce distortion.
  • Apply field map–based distortion correction in post-processing (available on most modern scanners).
  • For patients with metallic implants, use view-angle tilting or SEMAC (Slice Encoding for Metal Artifact Correction) techniques; these are available as proprietary sequences (e.g., MAVRIC, WARP).

In summary, a thorough understanding of the interplay between magnetization and susceptibility effects allows the MRI team to acquire high-quality T2*-weighted images that maximize diagnostic yield while minimizing artifact-related pitfalls.

Conclusion

The role of magnetization in T2*-weighted imaging is both fundamental and far-reaching. From the initial alignment of hydrogen spins in the magnetic field to the rapid dephasing caused by paramagnetic iron and deoxygenated blood, every step of signal evolution is governed by the magnetic properties of the tissue. Susceptibility effects, captured exquisitely by gradient-echo sequences, provide clinicians with an invaluable window into tissue composition and pathology that conventional spin-echo techniques cannot match. Advances such as SWI and QSM have transformed qualitative observation into quantitative biomarkers, enabling earlier diagnosis and more precise monitoring in neurological, cardiovascular, and systemic disorders. As MRI technology progresses toward higher field strengths and more sophisticated reconstruction algorithms, the importance of understanding magnetization and susceptibility will only grow. The principles outlined here equip imaging professionals with the knowledge needed to harness these effects for improved patient care.