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MRI pulse sequence

An MRI pulse sequence is a precisely timed series of radiofrequency (RF) pulses and magnetic field gradients applied during magnetic resonance imaging (MRI) to excite and manipulate the alignment and precession of hydrogen protons in tissue, thereby generating detectable signals that form images with specific contrast and spatial resolution.^[1]^[2]^[3] These sequences operate on the principles of nuclear magnetic resonance, where protons in a strong static magnetic field (B₀) align and precess at the Larmor frequency; an RF pulse, typically at 90° or 180° flip angles, perturbs this alignment, tipping magnetization into the transverse plane to induce a measurable electromotive force in receiver coils via relaxation processes (T1 longitudinal and T2 transverse).^[1]^[3] Key parameters include repetition time (TR), the interval between RF excitations, and echo time (TE), the duration from excitation to signal peak, which dictate image weighting: short TR and TE yield T1-weighted images emphasizing anatomical detail, while long TR and TE produce T2-weighted images highlighting pathology like edema.^[1] Common types include spin-echo and gradient-echo sequences, with advanced variants such as fast spin-echo and inversion recovery enhancing efficiency and specificity for clinical applications.^[1]^[2]

Historical Development

The foundations of MRI pulse sequences trace back to the discovery of nuclear magnetic resonance (NMR) by Isidor Isaac Rabi in 1938. In 1950, Erwin Hahn introduced the spin echo technique, which became fundamental for refocusing signals in MRI. The transition to imaging occurred in 1973 when Paul Lauterbur developed methods using magnetic field gradients to spatially encode NMR signals, enabling the first MRI images. Subsequent advancements in the 1980s and 1990s, including fast spin echo in 1986 by Jürgen Hennig and colleagues, significantly improved imaging speed and clinical utility.^[4]^[5] Pulse sequences are fundamental to MRI's versatility, enabling optimization of gradient waveforms for spatial encoding (slice selection, phase, and frequency directions) while adhering to hardware limits like maximum gradient strength and slew rate, thus minimizing artifacts and scan times in diagnostic protocols.^[2]

Introduction

Definition and Purpose

A magnetic resonance imaging (MRI) pulse sequence is defined as a predefined series of radiofrequency (RF) pulses, gradient magnetic fields, and specific timing intervals designed to manipulate the nuclear spins of hydrogen protons in tissues, thereby generating measurable transverse magnetization signals that form the basis for image reconstruction.^[3] These sequences control the excitation, evolution, and detection of the MR signal by precisely timing the application of RF pulses to tip spins away from the main magnetic field and using gradients to encode spatial information. The primary purpose of MRI pulse sequences is to enable selective control over image contrast, allowing differentiation of tissues based on intrinsic properties such as longitudinal (T1) and transverse (T2) relaxation times, proton density, diffusion coefficients, or blood flow characteristics, which is essential for identifying pathological changes in clinical diagnostics. By adjusting parameters like repetition time (TR), echo time (TE), and flip angles, sequences can emphasize specific contrast mechanisms—for instance, short TR and TE for T1-weighted images that highlight anatomical details, or longer values for T2-weighted images that reveal edema or inflammation. This versatility underpins MRI's non-invasive ability to provide high-resolution, multi-contrast imaging tailored to applications in neurology, oncology, and cardiology.^[1] In a typical pulse sequence workflow, the process begins with an excitation phase, often using a 90° RF pulse to rotate the net magnetization into the transverse plane, creating a detectable free induction decay (FID) signal. This is followed by a manipulation phase involving additional RF refocusing pulses or gradient applications to mitigate dephasing effects and enhance signal coherence, and concludes with a readout phase where frequency- and phase-encoding gradients are applied to acquire the spatial signal data during echo formation.^[6] The foundational concepts of MRI pulse sequences were first articulated in the 1970s by Paul Lauterbur, who demonstrated spatial encoding via gradients in 1973, and Peter Mansfield, who advanced rapid imaging techniques, contributions that earned them the 2003 Nobel Prize in Physiology or Medicine for establishing the principles of MRI.

Historical Development

The foundations of MRI pulse sequences trace back to the discovery of nuclear magnetic resonance (NMR) by Felix Bloch and Edward Purcell in 1946, who independently demonstrated the absorption of radiofrequency energy by atomic nuclei in magnetic fields, laying the groundwork for signal generation in both NMR and later MRI applications.^[7]^[8] Their work earned them the 1952 Nobel Prize in Physics, establishing the basic principles of spin manipulation through radiofrequency pulses that would evolve into MRI sequencing. In 1950, Erwin Hahn introduced the spin echo technique in NMR, using a 90°-180° radiofrequency pulse pair to refocus dephased spins and form an echo, which became the cornerstone for early MRI sequences by mitigating field inhomogeneities. This method was first applied to medical diagnostics in 1971 by Raymond Damadian, who used spin echo NMR to measure relaxation time differences in normal and cancerous tissues, proposing its use for tumor detection and marking the first link between NMR and medical diagnostics.^[9] Paul Lauterbur advanced this in 1973 by incorporating magnetic field gradients with spin echo signals to encode spatial information, enabling the first 2D NMR images and transforming pulse sequences into true imaging tools. The 1970s saw further innovations for efficiency, with Anil Kumar, Dieter Welti, and Richard Ernst introducing gradient echo sequences in 1975, which replaced the 180° refocusing pulse with a gradient reversal to generate echoes faster, facilitating Fourier transform-based 2D imaging and reducing scan times from hours to minutes. In 1977, Peter Mansfield developed echo-planar imaging (EPI), an ultrafast gradient echo variant that acquires an entire image plane in milliseconds via rapid gradient switching after a single excitation, enabling real-time applications and earning him the 2003 Nobel Prize in Physiology or Medicine shared with Lauterbur. Meanwhile, Edward Stejskal and James Tanner's 1965 pulsed gradient spin echo method for measuring diffusion in NMR was adapted clinically in the 1990s, incorporating bipolar gradients into sequences to quantify water molecule motion for detecting strokes and tumors.^[10] Clinical adoption accelerated with the U.S. Food and Drug Administration's approval of the first commercial MRI scanners in 1984, primarily using spin echo sequences for whole-body imaging at institutions like those employing Fonar systems. Contributions from David Hoult and Lauterbur in the late 1970s refined pulse design and sensitivity, optimizing radiofrequency coils and gradients for human-scale imaging while addressing power deposition limits. By the 2000s, standardization efforts through the DICOM protocol ensured interoperability of pulse sequence data across scanners, defining attributes like scanning sequence types (e.g., spin echo, gradient echo) to support consistent clinical workflows and multi-vendor integration. These developments were driven by demands for shorter scan times—from initial multi-hour acquisitions to seconds—and enhanced contrast for neurology and oncology, where sequences evolved to exploit T1/T2 relaxation and diffusion for tissue differentiation without ionizing radiation.^[5]

Fundamental Principles

RF Pulses, Gradients, and Echo Formation

Radiofrequency (RF) pulses are essential components in MRI pulse sequences, serving to excite and manipulate the magnetization of atomic nuclei, primarily hydrogen protons. The primary types include excitation pulses, typically with a 90° flip angle to tip the longitudinal magnetization into the transverse plane, and refocusing pulses, often 180° to reverse dephasing effects. The flip angle θ is determined by the gyromagnetic ratio γ of the nucleus, the amplitude of the RF magnetic field B₁, and the pulse duration τ, according to the relation θ = γ B₁ τ in the linear approximation for short pulses.^[11] To achieve spatial selectivity, such as slice selection, RF pulses are combined with magnetic field gradients; a frequency-modulated or shaped RF pulse applied during a slice-select gradient (usually G_z) excites spins only within a specific bandwidth corresponding to the desired slice thickness.^[12]^[13] Magnetic gradients play a crucial role in spatial encoding and signal manipulation within MRI pulse sequences. Three orthogonal gradients—G_x, G_y, and G_z—are used: G_z for slice selection during RF excitation, G_y for phase encoding to impart position-dependent phase shifts across one dimension, and G_x for frequency encoding to create position-dependent frequency variations during readout. These gradients induce dephasing by causing spins at different positions to precess at varying rates, which can be reversed through rephasing gradients of opposite polarity to form echoes. Dephasing and rephasing lobes in the gradient waveforms are precisely timed to control the spatial resolution and minimize artifacts.^[14]^[15] Echo formation is a fundamental process in MRI for generating detectable transverse magnetization signals after initial excitation. Following a 90° RF excitation pulse, the transverse magnetization produces a free induction decay (FID), a rapidly decaying signal due to T₂* relaxation and field inhomogeneities, which can be directly acquired but is limited by susceptibility effects. A spin echo is formed by applying a 180° refocusing RF pulse at time TE/2 after excitation; this pulse reverses the phase dispersion from static field inhomogeneities, refocusing the spins to produce a maximum signal at time TE, as originally demonstrated by Hahn in 1950. In contrast, a gradient echo is generated without a 180° pulse by applying a dephasing gradient lobe followed by a rephasing readout gradient of opposite polarity, refocusing the gradient-induced dephasing more rapidly but remaining sensitive to T₂* effects.^[16]^[17]^[18] Key timing parameters govern the evolution of the magnetization and the resulting signal. The echo time (TE) is the interval from the excitation RF pulse to the peak of the echo, influencing the degree of transverse relaxation, while the repetition time (TR) is the period between successive excitations, affecting longitudinal recovery. For gradient echo sequences, the signal intensity at time t follows S(t) ∝ M₀ sin(θ) e^{-t/T₂*}, where M₀ is the equilibrium magnetization, θ is the flip angle, and T₂* is the effective transverse relaxation time, highlighting the dependence on excitation efficiency and rapid decay.^[19]^[20] These elements—RF pulses, gradients, and echoes—combine to traverse k-space, the Fourier domain representation of the image, enabling spatial encoding for reconstruction. Gradients during phase and frequency encoding steps trace specific trajectories in k-space (e.g., Cartesian lines or spirals), sampling the spatial frequency components of the object; the inverse Fourier transform then reconstructs the spatial image from this data, as pioneered by Lauterbur in 1973 using projection reconstruction.^[21] This framework forms the prerequisite for all subsequent pulse sequences, allowing efficient filling of k-space to achieve high-resolution imaging.

Relaxation Processes and Contrast Mechanisms

In magnetic resonance imaging (MRI), contrast arises primarily from differences in tissue relaxation properties following radiofrequency excitation. Longitudinal relaxation, characterized by the time constant T1, describes the recovery of the net magnetization vector along the direction of the main magnetic field B0 after perturbation. This process involves energy transfer from excited spins to the surrounding lattice (spin-lattice relaxation), with the rate depending on molecular tumbling rates matching the Larmor frequency. The time evolution of the longitudinal magnetization Mz is given by

M_z(t) = M_0 \left(1 - e^{-t/T_1}\right),

where M0 is the equilibrium magnetization and t is time since excitation, assuming an initial Mz=0. T1 values typically range from 200-3000 ms in biological tissues at 1.5-3 T fields, longer in fluids like cerebrospinal fluid and shorter in fat-rich tissues due to efficient energy dissipation. Transverse relaxation, governed by the time constant T2, refers to the decay of the magnetization component perpendicular to B0 due to spin-spin interactions, leading to dephasing from local field variations induced by neighboring spins. The transverse magnetization Mxy decays as

M_{xy}(t) = M_0 e^{-t/T_2},

with T2 values generally shorter than T1 (e.g., 50-200 ms in gray matter at 3 T), reflecting irreversible loss of phase coherence. In practice, the observed decay is faster, described by T2*, which incorporates additional dephasing from magnetic field inhomogeneities and susceptibility effects:

\frac{1}{T_2^*} = \frac{1}{T_2} + \frac{1}{T_2'} ,

where T2' accounts for reversible extrinsic factors. T2 is the intrinsic tissue property, while T2* dominates in gradient-echo acquisitions.^[22] Beyond relaxation, proton density (PD) provides a baseline signal proportional to the number of mobile hydrogen nuclei per unit volume, contributing to contrast in low-relaxation-weighting regimes (e.g., PD-weighted images highlight differences in water content between tissues like muscle and fat). Diffusion, the Brownian motion of water molecules, modulates T2 contrast by enhancing dephasing in restricted environments, such as in edema or tumors, where reduced diffusivity alters local phase accrual. Perfusion introduces dynamic contrast from blood flow, affecting signal via delivery of oxygenated/deoxygenated hemoglobin or contrast agents, which influence T2* through susceptibility changes without direct relaxation modification.^[23]^[22] Contrast generation exploits these properties by adjusting repetition time (TR) and echo time (TE). Proton density weighting uses long TR (to allow full T1 recovery and minimize T1 effects) and short TE (to minimize T2 effects), emphasizing PD differences. T1 weighting employs short TR (to limit recovery and emphasize T1 differences) and short TE (to reduce T2 influence), highlighting tissues with short T1 (e.g., fat appears bright). T2 weighting requires long TR and long TE, accentuating tissues with long T2 (e.g., fluids appear bright). These parameters manipulate the steady-state signal as S ∝ PD · (1 - e^{-TR/T1}) · e^{-TE/T2}, enabling tissue differentiation based on biophysical variations.^[23] For quantitative imaging, accurate T1 and T2 mapping requires correction for radiofrequency field (B1) inhomogeneities, which cause flip-angle errors in variable flip-angle methods. B1 mapping techniques, such as the Bloch-Siegert shift, improve T1 precision by up to 20-30% in regions like the breast or brain, ensuring reliable relaxometry. Post-2010 advancements in multi-parametric mapping integrate T1, T2, and PD acquisitions for comprehensive tissue characterization, as in myocardial disease assessment, where T2 mapping detects inflammation with reproducibility errors below 5%.^[24]^[25]

Conventional Sequences

Spin Echo

The spin echo sequence is a foundational technique in magnetic resonance imaging (MRI), first demonstrated by Erwin Hahn in 1950 through experiments showing that a refocusing radiofrequency (RF) pulse could regenerate transverse magnetization after initial dephasing.^[26] In this sequence, a 90° excitation pulse tips the longitudinal magnetization into the transverse plane, initiating free induction decay (FID) that dephases due to spin-spin interactions and field inhomogeneities. A subsequent 180° refocusing pulse, applied at time TE/2 (where TE is the echo time), reverses the phase dispersion of spins, causing them to rephase and form a detectable echo signal at time TE. This refocusing mechanism distinguishes spin echo from simple FID, as it corrects for static magnetic field (B0) inhomogeneities and susceptibility-induced dephasing, yielding a signal that decays primarily according to the true transverse relaxation time T2 rather than the faster T2*. For multi-echo acquisitions, additional 180° pulses can generate successive echoes, allowing simultaneous collection of proton density and T2-weighted data within a single repetition. The sequence timing can be visualized as follows: following slice-selective gradients, the 90° RF pulse is applied, followed by phase-encoding and frequency-encoding gradients during the echo formation; the 180° pulse occurs midway, with readout gradients centered on the echo peak at TE. Key parameters include the repetition time (TR), the interval between successive 90° excitations, and TE. For T2-weighted imaging, a long TR (>2000 ms) minimizes T1 effects, while a long TE (80-120 ms) emphasizes T2 contrast; for proton density weighting, a long TR pairs with a short TE (<30 ms); and for T1 weighting, a shorter TR (500-800 ms) with a short TE enhances differences in longitudinal recovery. These parameters render the spin echo insensitive to B0 inhomogeneities, unlike gradient echo sequences, making it robust against artifacts from metal implants or air-tissue interfaces. As the gold standard for T2 contrast, the spin echo sequence excels in providing high-quality images with reduced susceptibility artifacts, enabling clear visualization of tissue pathology such as edema or inflammation. It is routinely employed in clinical neuroimaging and spine imaging to detect lesions, demyelination, and cerebrospinal fluid characteristics. To accelerate acquisition without sacrificing contrast, the fast spin echo (FSE) or turbo spin echo (TSE) variant—introduced by Hennig et al. in 1986 as rapid acquisition with relaxation enhancement (RARE)—uses multiple 180° refocusing pulses per TR to fill multiple lines of k-space, with echo train lengths (ETL) typically of 8-16 echoes. This reduces scan times by factors of 8-16 compared to conventional spin echo, making it suitable for high-resolution T2-weighted imaging in routine protocols. The signal intensity in a spin echo sequence follows the equation:

S = k \cdot \mathrm{PD} \left(1 - e^{-\mathrm{TR}/T_1}\right) e^{-\mathrm{TE}/T_2}

where S is the signal, k is a proportionality constant, PD is proton density, T_1 is the longitudinal relaxation time, and T_2 is the transverse relaxation time.^[20] This derives from the Bloch equations describing magnetization evolution. After the 90° pulse, the longitudinal magnetization M_z recovers toward equilibrium M_0 as M_z(t) = M_0 (1 - e^{-t/T_1}) during TR, contributing the (1 - e^{-\mathrm{TR}/T_1}) term; at the next excitation, this recovered M_z is tipped into the transverse plane. The transverse magnetization M_{xy} then dephases during the initial TE/2 due to T2 effects, but the 180° pulse inverts the phases, allowing rephasing over the subsequent TE/2, resulting in an echo amplitude attenuated solely by e^{-\mathrm{TE}/T_2}—unlike FID, where irreversible T2* decay (including inhomogeneity contributions) governs the signal without refocusing. This derivation assumes ideal pulses and neglects diffusion or higher-order effects, but it establishes the core contrast mechanisms reliant on T1 recovery and T2 decay.

Gradient Echo

Gradient echo (GRE) sequences in magnetic resonance imaging (MRI) employ a partial flip angle excitation, typically less than 90°, followed by gradient reversal to form an echo, without the use of a 180° refocusing radiofrequency (RF) pulse. This design contrasts with spin echo techniques by relying solely on magnetic field gradients for echo formation, making the sequence inherently sensitive to T2* relaxation effects, which include contributions from both true T2 decay and magnetic field inhomogeneities such as susceptibility-induced dephasing. As a result, GRE sequences produce images weighted by T2* contrast, which can highlight tissue differences due to local field variations but also introduce artifacts in regions with air-tissue interfaces or metal implants. Key parameters in GRE sequences include the echo time (TE), typically short at 2-5 ms to minimize T2* decay and maintain signal intensity, and the repetition time (TR), which can be varied from short (e.g., 5-20 ms) for rapid imaging to longer values for specific contrasts. The steady-state signal in GRE is governed by the Ernst equation, adjusted for T2* weighting:

S = \mathrm{PD} \cdot \sin\alpha \cdot \frac{1 - e^{-\mathrm{TR}/T_1}}{1 - \cos\alpha \cdot e^{-\mathrm{TR}/T_1}} \cdot e^{-\mathrm{TE}/T_2^*}

where PD is proton density, α is the flip angle, and T1 and T2* are the longitudinal and effective transverse relaxation times, respectively. This equation describes the transverse magnetization at steady state for low flip angles and short TR, optimizing signal through the Ernst angle (α_E = \arccos(e^{-\mathrm{TR}/T_1})), which balances excitation efficiency and T1 recovery.^[27] GRE variants include spoiled gradient echo (GRE) sequences, such as fast low-angle shot (FLASH), which apply RF spoiling or gradient crushing to eliminate residual transverse magnetization, yielding primarily T1-weighted images suitable for anatomical visualization. In contrast, balanced steady-state free precession (bSSFP) maintains a steady state without spoiling by balancing gradient moments, providing high signal-to-noise ratio (SNR) and T2/T1 contrast, often used in dynamic applications.^[28]^[27] The primary advantages of GRE sequences lie in their speed, enabling acquisition times of seconds per slice through short TR and TE, which facilitates 3D volumetric imaging and real-time applications like functional MRI. This rapidity stems from the absence of a 180° pulse, allowing higher duty cycles for RF transmission. However, limitations include heightened susceptibility to magnetic field inhomogeneities, leading to signal loss and distortions, particularly in bSSFP where off-resonance effects cause banding artifacts; recent advancements, such as interleaved radial linear combination bSSFP with partial dephasing, have mitigated these in cardiac imaging near implantable devices, improving cine quality for functional assessment.^[29]^[30]

Inversion-Based Sequences

Inversion Recovery

The inversion recovery (IR) sequence serves as a preparatory module in magnetic resonance imaging (MRI) to enhance T1 contrast by manipulating longitudinal magnetization prior to image acquisition. It begins with a 180° radiofrequency (RF) inversion pulse that flips the net magnetization vector along the longitudinal axis (Mz) from +M0 to -M0, where M0 represents the equilibrium magnetization. Following this, an inversion time (TI) delay allows partial recovery of Mz toward equilibrium through T1 relaxation, after which a standard 90° excitation pulse tips the recovered magnetization into the transverse plane for signal readout. The readout module can employ either a spin echo (SE) or gradient echo (GRE) acquisition to form the image, enabling flexibility in TE and TR parameters while preserving the T1-weighted preparation.^[31] The signal evolution in IR is governed by the T1 relaxation process during the TI period, yielding a longitudinal magnetization of:

M_z(\text{TI}) = M_0 \left(1 - 2 e^{-\text{TI}/T_1}\right)

where the factor of 2 accounts for the initial inversion. This results in a null point—where signal from a tissue is completely suppressed—when TI = T1 \ln(2) \approx 0.69 T1, allowing selective nulling based on tissue-specific T1 values. For T1-weighted imaging, an intermediate TI (typically 300-500 ms at 1.5 T) is selected to capture differences in recovery rates between tissues, producing high contrast where short-T1 tissues (e.g., fat and white matter) appear bright and long-T1 tissues (e.g., gray matter, fluids) appear relatively dark. Conversely, longer TI values (e.g., >2000 ms) can suppress signals from fluids with prolonged T1, though this is tuned carefully to avoid over-recovery. These parameters leverage the inherent T1 relaxation differences outlined in fundamental contrast mechanisms.^[31]^[32] IR offers distinct advantages in clinical MRI, particularly for enhancing T1 contrast to improve lesion detection in tissues with subtle relaxation differences, such as in neuroimaging or musculoskeletal applications. By providing superior gray-white matter differentiation compared to standard T1-weighted sequences, it aids in identifying pathologies like tumors or edema. Additionally, the inversion preparation module can be integrated as a magnetization transfer step in hybrid sequences, optimizing signal for subsequent acquisitions without altering core imaging parameters. Historically, IR was developed in the early 1980s as one of the foundational MRI techniques, initially for quantitative T1 mapping to characterize tissue properties accurately.^[31]^[32]

Variants for Specific Contrasts

Short tau inversion recovery (STIR) is an inversion recovery variant optimized for fat suppression by selecting an inversion time (TI) that nulls the fat signal, typically TI ≈ 140 ms at 1.5 T corresponding to the T1 relaxation time of fat (≈ 250 ms).^[33] This sequence employs a long repetition time (TR > 2000 ms) to allow recovery and is particularly effective at lower field strengths due to its field-independent fat suppression mechanism. Clinically, STIR is widely used to highlight edema and inflammation in musculoskeletal imaging, such as detecting bone marrow edema or soft tissue abnormalities, by enhancing the contrast of water-rich pathologies against suppressed fat background.^[34] Fluid-attenuated inversion recovery (FLAIR) adapts inversion recovery with a long TI (2000-2500 ms) to suppress cerebrospinal fluid (CSF) signal, which has a prolonged T1, while using an extended TR (> 8000 ms) to maintain T2-weighted contrast in brain tissue.^[35] This results in dark CSF on images, preventing its hyperintensity from obscuring adjacent lesions, and is essential for neuroimaging protocols. FLAIR excels in detecting periventricular and subcortical brain lesions, such as those in multiple sclerosis or stroke, by improving lesion conspicuity without CSF interference.^[36] Magnetization-prepared rapid acquisition of gradient echoes (MP-RAGE) combines inversion recovery preparation with a rapid gradient echo readout to produce high-resolution three-dimensional T1-weighted images, typically using TI ≈ 900 ms and a low flip angle (α ≈ 10°) to optimize gray-white matter contrast.^[37] The inversion pulse enhances T1 weighting, followed by segmented k-space acquisition for efficiency, enabling isotropic voxels for detailed structural brain imaging. It is a standard for volumetric analysis in neuroimaging, supporting applications like cortical thickness measurement and atrophy assessment in neurodegenerative diseases.^[38] Double inversion recovery extends the technique by applying two 180° inversion pulses to suppress signals from two distinct tissues simultaneously, such as CSF (TI1 ≈ 2000-3000 ms) and white matter (TI2 ≈ 450 ms), on a T2-weighted background.^[39] This dual suppression improves the detection of gray matter lesions, particularly in multiple sclerosis, by reducing background noise from normal tissues. Clinical protocols often incorporate it for spinal cord or infratentorial imaging, though artifacts from imperfect inversion—due to B1 inhomogeneities or motion—can lead to incomplete suppression and require careful parameter adjustment.^[40]

Motion and Diffusion Sequences

Diffusion-Weighted Imaging

Diffusion-weighted imaging (DWI) is an MRI technique that sensitizes the signal to the random Brownian motion of water molecules in tissues by applying pairs of pulsed magnetic field gradients. This method exploits the fact that diffusion is restricted in certain pathological states, such as cellular swelling, allowing for the detection of microstructural changes. The core principle relies on the Stejskal-Tanner equation, which describes the signal attenuation due to diffusion:

S = S_0 e^{-b D}

where S is the observed signal intensity, S_0 is the signal without diffusion weighting, D is the apparent diffusion coefficient, and b is the b-value that quantifies the degree of diffusion sensitization.^[10] The b-value is calculated as b = \gamma^2 \delta^2 g^2 (\Delta - \delta/3), with \gamma as the gyromagnetic ratio, \delta as the gradient duration, g as the gradient amplitude, and \Delta as the time interval between the leading edges of the gradients; typical b-values range from 0 to 1000 s/mm² to balance sensitivity and signal-to-noise ratio.^[41] The standard DWI sequence is based on a spin-echo framework, where diffusion-sensitizing gradients are applied before and after the 180° refocusing pulse to encode motion-induced phase shifts. A single-shot echo-planar imaging (EPI) readout is commonly used immediately following the spin echo to acquire the entire image plane rapidly, minimizing motion artifacts and enabling T2*-weighted contrast.^[41] Key parameters include an echo time (TE) of 60-100 ms to accommodate gradient timing while preserving signal, and acquisition of images at multiple b-values (e.g., b=0 and b=1000 s/mm²) to compute apparent diffusion coefficient (ADC) maps via pixel-wise fitting of the Stejskal-Tanner equation. For isotropic diffusion measurement, gradients are applied in three orthogonal directions and averaged, whereas directional sensitivity is achieved by varying gradient orientations to probe tissue anisotropy.^[41] Clinically, DWI excels in detecting acute ischemic stroke, where cytotoxic edema restricts water diffusion, yielding hyperintense signals on trace images and low ADC values in the infarct core within minutes of onset. Diffusion tensor imaging (DTI), an extension of DWI, acquires data in at least six non-collinear directions to model diffusion as a tensor, enabling quantification of fractional anisotropy (FA)—a scalar measure (0-1) of directional preference in water motion that highlights organized structures like white matter tracts for fiber tracking in neuroimaging. Common artifacts in DWI include distortions from eddy currents induced by rapid gradient switching, which cause geometric warping, and bulk motion leading to signal loss or ghosting. These are mitigated through pre-emphasis calibration of gradients for eddy current compensation and navigator echoes or prospective motion correction to track and adjust for patient movement during acquisition.^[42]

Perfusion-Weighted Imaging

Perfusion-weighted imaging (PWI) in MRI quantifies tissue perfusion by measuring blood flow and volume using either exogenous contrast agents or endogenous blood water as tracers. Exogenous techniques rely on gadolinium-based contrast agents (GBCAs) injected as a bolus to track hemodynamic changes, while endogenous methods magnetically label arterial blood without contrast. These sequences are essential for assessing microvascular dynamics in clinical settings, providing parameters such as cerebral blood flow (CBF) and cerebral blood volume (CBV).^[43] Dynamic susceptibility contrast (DSC)-MRI is a primary exogenous method that uses T2*-weighted gradient-echo echo-planar imaging (GRE-EPI) to detect transient signal drops caused by the susceptibility effects of a GBCA bolus passing through capillaries. The signal intensity decreases due to local magnetic field inhomogeneities induced by the contrast agent, allowing measurement of the first-pass bolus dynamics. The transverse relaxation rate change is modeled as

R_2^*(t) = R_2^*_0 + \Delta R_2^* [\text{Gd}(t)]

, where

R_2^*_0

is the baseline rate, \Delta R_2^* is the susceptibility relaxivity, and [\text{Gd}(t)] is the gadolinium concentration over time; this equation enables derivation of concentration-time curves from signal changes. Dynamic contrast-enhanced (DCE)-MRI complements DSC by employing T1-weighted GRE sequences to evaluate vessel permeability and extravasation, tracking the T1-shortening effects of GBCA leakage into the extravascular space for parameters like the volume transfer constant K^\text{trans}. Perfusion quantification in both requires an arterial input function (AIF) derived from major arteries, followed by deconvolution techniques such as singular value decomposition (SVD) to compute CBF from the impulse response function; CBV is obtained by integrating the tissue concentration-time curve, often normalized to gray matter values.^[44]^[45]^[46]^[47]^[48]^[49]^[50] Arterial spin labeling (ASL) provides a non-invasive endogenous alternative by inverting arterial blood spins upstream of the imaging slice to create a diffusible tracer from water protons. Flow-sensitive alternating inversion recovery (FAIR), a pulsed ASL variant, applies slice-selective and non-selective inversion pulses to label blood selectively, while pseudo-continuous ASL (pCASL) uses a train of rapid RF pulses with gradients for efficient, multi-slice labeling and higher signal-to-noise ratio. The quantitative imaging of perfusion using a single subtraction (QUASAR) variant enhances accuracy by incorporating saturation pulses to control bolus length and multi-inversion time sampling for robust CBF estimation. Recent advancements in the 2020s include AI-accelerated ASL, such as deep learning networks that reduce required post-labeling delays by over 60% while maintaining quantitative CBF and arterial transit time measurements, aiding standardization efforts.^[43]^[51]^[52]^[53]^[54] Clinically, PWI aids in glioma grading by leveraging elevated relative CBV (rCBV > 1.75) in high-grade tumors to predict aggressiveness and guide biopsy targeting. In acute stroke, it identifies the ischemic penumbra—hypoperfused but viable tissue—through perfusion-diffusion mismatch, where delays in time-to-peak (>6 seconds) indicate salvageable regions for intravenous thrombolysis up to 9 hours post-onset in select patients. These applications underscore PWI's role in distinguishing tumor recurrence from pseudoprogression and optimizing stroke interventions.^[55]^[56]^[57]^[58]

Functional and Vascular Sequences

Functional MRI

Functional magnetic resonance imaging (fMRI) is a neuroimaging technique that maps brain activity by detecting changes in blood oxygenation levels, primarily through the blood oxygenation level-dependent (BOLD) contrast mechanism. The BOLD effect arises from the paramagnetic properties of deoxyhemoglobin, which increases magnetic susceptibility and accelerates T2* decay in regions with lower oxygenation, leading to signal loss in gradient-echo sequences. During neural activation, cerebral blood flow rises disproportionately to oxygen consumption, reducing deoxyhemoglobin concentration and thereby increasing the T2*-weighted signal in active brain areas.^[59] This non-invasive method was first demonstrated in humans using sensory stimulation tasks, enabling real-time mapping of cortical responses without exogenous contrast agents.^[59] The standard pulse sequence for BOLD fMRI is echo-planar imaging (EPI) combined with gradient echo (GRE), optimized for rapid whole-brain coverage and sensitivity to T2* changes.^[60] Key parameters include an echo time (TE) of approximately 30-40 ms to maximize BOLD contrast near the T2* peak, and a repetition time (TR) of 2-3 seconds for adequate temporal sampling of hemodynamic responses.^[61] Typical spatial parameters feature a matrix size of 64x64 to 128x128 voxels and slice thickness of 3 mm, balancing resolution with signal-to-noise ratio and scan time.^[62] These settings allow acquisition of multiple slices covering the brain in under 3 seconds, essential for capturing dynamic activity patterns. fMRI experiments employ two main paradigms: task-based designs, such as block or event-related protocols where subjects perform alternating periods of stimulation and rest, and resting-state paradigms that analyze spontaneous low-frequency fluctuations in the absence of tasks.^[60] Data preprocessing is critical and includes steps like motion correction to realign images across time points, spatial normalization to a standard template, and smoothing to enhance signal detection.^[63] Statistical analysis typically uses the general linear model (GLM) to fit voxel-wise time series against expected hemodynamic response functions, generating activation maps via t-tests or F-statistics on parameter estimates.^[64] Recent advances have improved fMRI efficiency, particularly through multiband EPI, which simultaneously excites multiple slices to achieve higher acceleration factors (e.g., 4-8x) and sub-second whole-brain temporal resolution in the 2020s. Emerging sequences like zero echo time (ZTE) fMRI provide artifact-free and quiet imaging alternatives, reducing susceptibility distortions and acoustic noise for better brainstem and spinal cord mapping.^[65] At ultra-high fields like 7T, enhanced BOLD sensitivity and spatial resolution enable finer mapping of subcortical structures and laminar cortical activity, though challenges such as increased susceptibility artifacts persist. Further progress at 9.4T using 3D stack-of-spirals readouts has demonstrated even higher resolution for preclinical and advanced human studies as of 2025.^[66]

Magnetic Resonance Angiography

Magnetic Resonance Angiography (MRA) encompasses a group of MRI techniques designed for non-invasive visualization of blood vessels, primarily relying on flow-induced signal alterations rather than exogenous contrast agents. These methods exploit the differences in magnetization between stationary tissues and flowing blood to generate angiographic images, enabling assessment of vascular anatomy and pathology without catheterization.^[67] Time-of-Flight (TOF) MRA is a cornerstone non-contrast technique that utilizes gradient echo sequences with short repetition time (TR) and echo time (TE) to achieve flow-related enhancement. In this approach, repeated radiofrequency pulses saturate the longitudinal magnetization of stationary spins within the imaging volume, while inflowing blood from outside the slice remains unsaturated and yields high signal intensity upon excitation. This contrast mechanism is particularly effective for arteries with steady inflow perpendicular to the imaging plane.^[68] TOF-MRA can be implemented in two-dimensional (2D) mode, acquiring thin slices sequentially for better depiction of tortuous vessels, or three-dimensional (3D) mode using slab excitation for isotropic high-resolution imaging of vessel segments, though 3D acquisitions are more prone to intravoxel dephasing in longer flow paths.^[67] Key parameters for TOF-MRA include TR values under 50 ms to minimize saturation of inflowing spins, TE as short as possible (typically 2-7 ms) to reduce T2* decay and susceptibility artifacts, and flip angles of 20-30° to balance background suppression with blood signal optimization. Spatial presaturation bands are routinely applied upstream or downstream of the imaging slab to suppress signals from veins or unwanted arterial branches, enhancing arterial specificity. These settings allow for rapid acquisitions, often under 5-10 minutes for intracranial studies, while maintaining vessel-to-background contrast.^[69]^[70] Phase Contrast (PC) MRA complements TOF by quantifying blood flow velocity and direction through phase differences induced by motion-sensitive gradients, making it suitable for both qualitative angiography and hemodynamic assessment. Bipolar velocity-encoding gradients are applied along one or more directions, creating a first-order moment that imparts a phase shift proportional to spin velocity; the phase φ is calculated as \phi = \gamma \int \mathbf{G} \cdot \mathbf{r} \, dt, where \gamma is the gyromagnetic ratio, \mathbf{G} is the gradient vector, and \mathbf{r} is the position of the moving spin. Magnitude and phase images are reconstructed from acquisitions with and without encoding, with the phase map directly reflecting velocity.^[71] To avoid phase wrapping (aliasing), the velocity encoding parameter (VENC) is selected to match the anticipated peak flow velocity, typically 50-150 cm/s for cerebral arteries or up to 200 cm/s for larger vessels; velocities exceeding VENC/2 cause aliasing, necessitating higher VENC values at the cost of signal-to-noise ratio for slower flows. PC-MRA supports multi-directional encoding for comprehensive flow mapping but requires longer scan times (10-20 minutes) due to multiple phase-encoding steps.^[72]^[71] Clinically, TOF-MRA excels in screening for carotid artery stenosis, where it accurately depicts luminal narrowing with sensitivity over 90% for high-grade lesions, and intracranial aneurysms, providing non-invasive detection of saccular dilatations as small as 3 mm. PC-MRA adds value in evaluating flow dynamics across stenoses or in aneurysm necks, aiding in rupture risk assessment by quantifying turbulence or jet velocities. Both techniques are preferred for pediatric and renal-impaired patients due to the absence of iodinated contrast risks.^[73]^[72] Despite their advantages, TOF-MRA suffers from signal voids in regions of slow, turbulent, or in-plane flow, potentially leading to overestimation of stenosis or underdetection of aneurysms with complex geometries. PC-MRA is limited by sensitivity to higher-order motion (e.g., acceleration) and reduced accuracy in non-laminar flows, such as those in severe stenoses or aneurysms, where phase dispersion occurs. Scan times and susceptibility to patient motion remain challenges for both.^[67]^[74] Recent advances in MRA include applications at ultra-high fields like 7T, which enhance resolution for imaging small intracranial vessels and perforators, improving detection of microvascular pathologies such as vasculitis, as of 2025.^[75] For cases where flow artifacts compromise non-contrast MRA, contrast-enhanced MRA using gadolinium chelates with T1-weighted gradient echo sequences provides robust vessel opacification by shortening blood T1 relaxation, though it introduces risks like nephrogenic systemic fibrosis in vulnerable populations.^[67]

Advanced Contrast Sequences

Susceptibility Weighted Imaging

Susceptibility weighted imaging (SWI) is a gradient-echo-based magnetic resonance imaging technique that exploits differences in magnetic susceptibility among tissues to generate high-contrast images, particularly sensitive to paramagnetic substances such as deoxyhemoglobin, iron deposits, and calcium. The method combines magnitude and phase information from a fully flow-compensated, high-resolution three-dimensional gradient-echo sequence to amplify susceptibility-induced effects, enabling visualization of small venous structures, hemorrhages, and mineral accumulations that are subtle in conventional imaging. Unlike standard T2*-weighted gradient-echo sequences, which primarily rely on magnitude signal decay due to T2* relaxation, SWI incorporates post-processing steps to selectively enhance phase variations arising from local field inhomogeneities. The SWI sequence typically employs a three-dimensional spoiled gradient-recalled echo acquisition with echo times (TE) of 20-40 ms to accumulate sufficient phase accrual from susceptibility sources, repetition times (TR) of 25-50 ms, and flip angles of 12-20° (often 15°) to optimize signal-to-noise ratio while maintaining steady-state conditions. High spatial resolution, such as 0.5 mm isotropic voxels, is achieved through thin slices and matrix sizes up to 512×512, often at field strengths of 3T or higher for enhanced susceptibility contrast. Multiple echoes may be acquired in advanced implementations to improve phase reliability and enable multi-echo processing for better background suppression. The core post-processing involves creating a phase mask from the filtered phase image, obtained via homodyne high-pass filtering. This is done by dividing the original complex image by its low-pass filtered version (using a slice-by-slice 2D Hanning window of size e.g., 64×64), and taking the phase of the resulting high-pass filtered complex image to obtain the filtered phase \phi_f(\mathbf{r}). The phase mask m(\mathbf{r}) is then generated from the negative (paramagnetic) portion of the filtered phase, assuming a convention where negative phase indicates increased susceptibility:

m(\mathbf{r}) = \begin{cases} \frac{\pi + \phi_f(\mathbf{r})}{\pi} & -\pi \leq \phi_f(\mathbf{r}) \leq 0 \\ 1 & 0 < \phi_f(\mathbf{r}) \leq \pi \end{cases}

The final SWI image is the magnitude multiplied by the mask raised to a power (e.g., m^4) for contrast enhancement. This processing distinguishes SWI from basic gradient-echo by providing susceptibility-specific amplification without altering the acquisition.^[76] SWI finds primary applications in neuroimaging for detecting cerebral microbleeds, venous structures in acute stroke, and lesions in multiple sclerosis, where it reveals perivenular iron deposition and demyelination-related susceptibility changes with superior sensitivity compared to T2*-weighted imaging. In trauma and neurodegenerative diseases, it identifies hemosiderin from prior hemorrhages and abnormal iron accumulation in basal ganglia. An extension, quantitative susceptibility mapping (QSM), derives absolute susceptibility values in parts per billion by solving the ill-posed inverse problem of dipole field deconvolution, often using morphology-enabled regularization to handle streaking artifacts. Recent integrations of artificial intelligence, such as deep learning-based acceleration and regularization in QSM pipelines, have improved reconstruction speed and accuracy for clinical deep brain nuclei evaluation as of 2025.^[77]^[78]

Magnetization Transfer and Fat Suppression

Magnetization transfer (MT) imaging enhances contrast by exploiting the exchange of magnetization between free water protons and those bound to macromolecules, such as proteins and lipids in tissues. This technique employs an off-resonance radiofrequency (RF) pulse that selectively saturates the protons in the immobile macromolecular pool, which have a broader resonance linewidth due to their restricted motion. The saturation indirectly reduces the signal from the free water pool through cross-relaxation and chemical exchange, without directly exciting the free water protons at their on-resonance frequency.^[79]^[80] The magnetization transfer ratio (MTR), a common quantitative measure, is calculated as MTR = (M₀ - M_sat) / M₀, where M₀ is the signal intensity without the MT pulse and M_sat is the signal with the saturation pulse applied. This ratio reflects the efficiency of magnetization exchange and is typically expressed as a percentage. In practice, the MT pulse—a series of high-power, off-resonance RF pulses—is added as a preparation module to standard gradient echo (GRE) or spin echo (SE) sequences, allowing integration into routine imaging protocols without significantly extending scan times.^[81]^[82] MT imaging finds key applications in quantifying myelin content and detecting pathological changes, such as in multiple sclerosis plaques where reduced MTR correlates with demyelination. For instance, lower MTR values in white matter lesions indicate disrupted macromolecular integrity, aiding in early plaque identification and monitoring disease progression. Clinically, it improves visualization of brain tissue microstructure, particularly for myelin-sensitive contrasts in neurological assessments.^[83]^[84] Fat suppression techniques are essential in MRI to eliminate the bright signal from adipose tissue, which can otherwise obscure pathology or mimic lesions due to its short T1 relaxation time. Several methods achieve this by targeting the chemical shift difference between fat and water protons, approximately 3.5 ppm at 1.5 T, allowing selective manipulation of fat magnetization. Chemical shift selective (CHESS) saturation uses a frequency-selective 90° RF pulse tuned to the fat resonance (offset by -3.5 ppm from water), followed by a spoiler gradient to dephase the excited fat spins, and sometimes preceded by a 180° pulse for better homogeneity in variants like the 90°-180°-90° scheme. This method provides uniform suppression in homogeneous fields but is sensitive to B₀ and B₁ inhomogeneities, particularly at higher field strengths. Short tau inversion recovery (STIR), an inversion recovery-based approach, applies a non-selective 180° inversion pulse followed by a short inversion time (TI ≈ 140-170 ms at 1.5 T) to null the fat signal, as detailed in inversion recovery sequences. It offers robust suppression independent of field strength but suppresses all short T1 tissues, limiting its use with contrast agents. Spectral adiabatic inversion recovery (SPAIR) combines spectral selectivity with adiabatic inversion pulses, which are insensitive to B₁ variations, to invert only fat protons before allowing recovery to null the signal at a specific TI. This technique excels in regions with field inhomogeneities, such as the extremities, providing more homogeneous suppression than CHESS. The Dixon method, a phase-based water-fat separation technique, acquires images at echo times where fat and water signals are in-phase and opposed-phase, modeling the complex signal as S = |W + F e^{iφ}|, where W and F are water and fat magnitudes, and φ encodes the chemical shift phase evolution (φ = 2π Δf TE + ψ, with Δf ≈ 220 Hz at 1.5 T). Iterative decomposition solves for separate water and fat images, yielding robust suppression even in inhomogeneous fields and enabling fat fraction quantification. Seminal work by Dixon introduced this opposed-phase approach, with modern multi-point variants improving accuracy.^[85] In clinical practice, these fat suppression methods are vital in orthopedics for detecting bone marrow edema and soft-tissue injuries, where CHESS or SPAIR enhances contrast in T2-weighted images of the musculoskeletal system. In breast imaging, Dixon and SPAIR techniques improve lesion conspicuity by reducing fat artifacts, facilitating accurate assessment of enhancing tumors in dynamic contrast-enhanced protocols.^[86]

Specialized and Emerging Sequences

Neuromelanin Imaging

Neuromelanin imaging utilizes T1-weighted MRI sequences to detect neuromelanin, a dark pigment accumulated in catecholaminergic neurons of the substantia nigra pars compacta (SNc) and locus coeruleus (LC). The paramagnetism arising from neuromelanin-iron complexes shortens T1 relaxation times, producing hyperintense signals in these midbrain structures on T1-weighted images without requiring exogenous contrast agents. This endogenous contrast arises specifically from the melanin-bound iron, enabling visualization of regions vulnerable to degeneration in Parkinson's disease (PD).^[87]^[88] Common pulse sequences for neuromelanin imaging include 2D T1-weighted turbo spin-echo (TSE) and 3D gradient-recalled echo (GRE) with magnetization transfer (MT) preparation, typically acquired at 3T. For 2D TSE, standard parameters are repetition time (TR) of 600 ms, echo time (TE) of 10 ms, echo train length of 8, in-plane resolution of 0.6 × 0.6 mm, and slice thickness of 2-3 mm over 15-20 axial slices centered on the midbrain. 3D GRE sequences enhance volume coverage and incorporate MT pulses (e.g., Gaussian-shaped with flip angle 500°, duration 10 ms) to boost contrast, though they may use shorter TR (e.g., 55 ms) and TE (e.g., 2.2 ms) for efficiency. These parameters optimize the paramagnetic T1 shortening while minimizing artifacts from surrounding tissues.^[89]^[90]^[91] Clinically, neuromelanin imaging supports early PD diagnosis by quantifying signal loss and atrophy in the SNc and LC, markers of dopaminergic and noradrenergic neuron degeneration. Signal intensity is often assessed via contrast ratios (CR) or contrast-to-noise ratios (CNR) normalized to reference regions like white matter or the pontine tegmentum; in PD, SNc CR decreases by approximately 20-30% compared to controls, correlating with disease severity. Volumetric analysis of the SNc further aids differentiation from essential tremor, with thresholds below 440 mm³ indicating PD. Atrophy measurement tracks progression, providing a noninvasive biomarker for monitoring therapeutic responses.^[87]^[90]^[92] Development of neuromelanin-sensitive sequences accelerated in the 2000s, building on postmortem validations of neuromelanin's MRI properties. Seminal in vivo demonstrations at 3T were reported by Sasaki et al. in 2006 using TSE sequences, establishing reliable SNc visualization. Subsequent optimizations in the 2010s focused on reproducibility, MT weighting, and quantitative metrics like signal ratios. Recent 2020s advancements at 7T exploit higher field strength for improved signal-to-noise ratio and submillimeter resolution (voxel volumes ~0.24 mm³), enhancing LC depiction and diagnostic performance—achieving 100% sensitivity and 96.8% specificity for SNc volumetry in PD—while maintaining similar T1-based principles.^[93]^[94]

T1 Rho Imaging

T1ρ imaging, also known as spin-lattice relaxation in the rotating frame, measures the relaxation of longitudinal magnetization in the presence of a spin-lock radiofrequency (RF) field, providing sensitivity to low-frequency molecular motions in tissues.^[95] The principle relies on applying a continuous or pulsed spin-lock field (B₁) along the transverse plane after excitation, which locks the magnetization vector and suppresses dipolar interactions at low locking frequencies (typically 200–1000 Hz), making T1ρ relaxation time approximate T2 at these low frequencies while being particularly sensitive to chemical exchange processes between water protons and macromolecules like proteoglycans.^[95] This contrasts with standard T1 and T2 relaxations by probing interactions on the order of the spin-lock frequency, enabling detection of subtle biochemical changes not visible in conventional MRI.^[96] The pulse sequence begins with a 90° excitation pulse to tip magnetization into the transverse plane, followed by a spin-lock module consisting of continuous wave (CW) RF pulses or composite pulses such as self-compensated spin-lock (SLR) trains to maintain the lock for varying durations (τ, or spin-lock time TSL).^[95] Multiple images are acquired at different τ values, often using multi-echo acquisition for signal averaging and fitting, with the signal intensity decaying mono-exponentially as S(\tau) = S_0 e^{-\tau / T_{1\rho}}, where S(\tau) is the observed signal, S_0 is the equilibrium signal, and T_{1\rho} is the relaxation time constant obtained via voxel-wise curve fitting.^[95] For improved efficiency in 3D imaging, balanced steady-state free precession (bSSFP)-based sequences, such as T1ρ-prepared bSSFP, incorporate centric k-space ordering and transient signal filtering to achieve rapid acquisitions (under 10 minutes for knee coverage) while preserving T1ρ contrast and reducing specific absorption rate (SAR) concerns.^[97] T1ρ dispersion curves, plotting T1ρ values against varying spin-lock frequencies, reveal tissue-specific low-frequency relaxation behaviors influenced by macromolecular content, aiding in the characterization of pathological changes.^[95] Primary applications include quantifying early cartilage degeneration in osteoarthritis (OA), where elevated T1ρ values (e.g., 55.4 ± 26.0 ms in advanced knee OA) correlate with proteoglycan loss and outperform T2 in sensitivity; intervertebral disc degeneration, detecting annulus fibrosus changes with steeper T1ρ gradients (e.g., -3.02 to -4.56); and amyloid detection in Alzheimer's disease, with increased hippocampal T1ρ associated with cognitive impairment and amyloid burden.^[95]00950-0/fulltext)^[98] Currently in transition from experimental to clinical use, T1ρ imaging faces challenges like SAR limitations and scan time but benefits from acceleration techniques like parallel imaging, positioning it as a promising biomarker for early musculoskeletal and neurodegenerative pathologies without exogenous contrast.^[95]^[99]

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Moreover, TOF-MRA is prone to saturation of slowly flowing blood protons, resulting in signal loss within a severely compromised lumen or distal small arteries ...
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[PDF] Susceptibility-Weighted Imaging: Technical Aspects and Clinical ...
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Morphology Enabled Dipole Inversion for Quantitative Susceptibility ...
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Validation of deep-learning accelerated quantitative susceptibility ...
Jan 22, 2025 · Abstract. Purpose. To test the feasibility and consistency of a deep-learning (DL) accelerated QSM method for deep brain nuclei evaluation.
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[PDF] Magnetization transfer in MRI: a review
As shown in. Fig. 1, it is possible to saturate the macromolecular spins preferentially using an off-resonance radio frequency pulse. The macromolecular spins ...
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Magnetization Transfer - Questions and Answers in MRI
This off-resonance pulse has sufficient power to saturate protons in the immobile pool without directly affecting those in free water. Following this MT ...
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[PDF] Magnetization Transfer Basics - ISMRM
The MT effect causes additional magnetization decrease (solid line MMT). The magnetization transfer ratio, MTR is a sum of Mdir and MMT contributions.
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Magnetization Transfer - Quantitative MRI mOOC - qMRLab
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Myelin Measurement Using Quantitative Magnetic Resonance Imaging
A widely used myelin imaging method is based on magnetization transfer (MT), which is a physical process by which macromolecules and their closely associated ...
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Quantitative magnetization transfer imaging in relapsing-remitting ...
Apr 4, 2022 · Myelin-sensitive MRI such as magnetization transfer imaging has been widely used in multiple sclerosis. The influence of methodology and ...
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Dixon techniques for water and fat imaging - Ma - Wiley Online Library
Sep 5, 2008 · The primary objective of the Dixon techniques is to determine W and F from an acquired image or images that are represented by S in Equation [1] ...Missing: exp( | Show results with:exp(
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Fat suppression techniques in breast magnetic resonance imaging
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Neuromelanin detection by magnetic resonance imaging (MRI) and ...
Apr 10, 2018 · We discuss how MRI detects NM and how this approach might be improved. We suggest that MRI of NM can be used to confirm PD diagnosis and monitor disease ...
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[PDF] Normative Values of Neuromelanin‐Sensitive MRI Signal in Older ...
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Neuromelanin-sensitive MRI as a promising biomarker of ...
MRI studies with NM-MRI sequences have consistently demonstrated the presence of a significant reduction of NM-related signal in the SN (SNc-VTA) and LC in ...
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Reproducibility assessment of neuromelanin-sensitive magnetic ...
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Evaluation of Parkinson Disease and Alzheimer Disease with the ...
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T1ρ magnetic resonance: basic physics principles and applications ...
Dec 6, 2015 · This paper reviews the T 1ρ imaging's basic physic principles, as well as its application for cartilage imaging and intervertebral disc imaging.
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T1ρ MR Imaging of Human Musculoskeletal System - PMC - NIH
In the following sections of this article, we will review the basic principles of T1ρ MR imaging, implementation of different T1ρ pulse sequences, biochemical ...
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T1ρ-Prepared Balanced Gradient Echo for Rapid 3D T1ρ MRI - NIH
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Comparison of T1rho MRI, Glucose Metabolism and Amyloid Burden ...
Higher T1rho values were associated with a clinical diagnosis of cognitive impairment (i.e.,MCI or AD diagnosis), poorer neuropsychological test performance ( ...
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MR-Imaging in Osteoarthritis: Current Standard of Practice and ... - NIH
Aug 3, 2023 · T1-Rho Mapping. T1rho relaxation is also known as spin-lattice relaxation in the r(h)otational frame. In contrast to conventional T1 ...2. Osteoarthritis--A Whole... · 3.4. Fat Suppression · 6.3. T2 Mapping