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Short-T2 MRI: Principles and recent advances
Journal Pre-Proof Short-T2 MRI: Principles and recent advances Markus Weiger, Klaas P. Pruessmann PII: DOI: Reference:
S0079-6565(19)30033-0 https://doi.org/10.1016/j.pnmrs.2019.07.001 JPNMRS 1483
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Progress in Nuclear Magnetic Resonance Spectroscopy
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Please cite this article as: M. Weiger, K.P. Pruessmann, Short-T2 MRI: Principles and recent advances, Progress in Nuclear Magnetic Resonance Spectroscopy (2019), doi: https://doi.org/10.1016/j.pnmrs.2019.07.001
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JOURNAL PRE-PROOF Short-T2 MRI: Principles and recent advances Markus Weiger*,[email protected], Klaas P. Pruessmann
Institute for Biomedical Engineering, ETH Zurich and University of Zurich, Zurich, Switzerland *Corresponding
author at: Institute for Biomedical Engineering, ETH Zurich and University of
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Zurich, Gloriastrasse 35, CH-8092 Zurich, Switzerland.
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Graphical abstract Highlights
Overarching concepts structure the variety of short-T2 MRI techniques
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Relaxation in a nutshell
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Detailed description of ZTE imaging and its variants
Dedicated hardware requirements for all scanner components
Abstract
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Comprehensive literature overview including T2 and T2* values of potential imaging targets
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Among current modalities of biomedical and diagnostic imaging, MRI stands out by virtue of its versatile contrast obtained without ionizing radiation. However, in various cases, e.g., water protons in tissues such as bone, tendon, and lung, MRI performance is limited by the rapid decay of resonance
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signals associated with short transverse relaxation times T2 or T2*. Efforts to address this shortcoming have led to a variety of specialized short-T2 techniques. Recent progress in this field expands the
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choice of methods and prompts fresh considerations with regard to instrumentation, data acquisition, and signal processing. In this review, the current status of short-T2 MRI is surveyed. In an attempt to
structure the growing range of techniques, the presentation highlights overarching concepts and basic methodological options. The most frequently used approaches are described in detail, including acquisition strategies, image reconstruction, hardware requirements, means of introducing contrast, sources of artifacts, limitations, and applications. Keywords: Relaxation; Zero echo time (ZTE); Ultra-short echo time (UTE); Single point imaging (SPI); Solid tissues
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
1 Introduction Magnetic resonance imaging (MRI) is a well-established and widely used modality in medical diagnostics due to its ability to depict tissues in three dimensions, with versatile contrast, and without ionizing radiation. It is based on spatial encoding and detection of macroscopic transverse nuclear magnetization. The latter is of limited lifetime and usually decays approximately exponentially with
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time constants T2 or T2*, depending on the exact nature of the decay. Nuclei of molecules in liquids, in
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particular water protons in most soft tissues, are characterized by relatively long T2 in the range of tens of milliseconds to seconds, which readily permit extensive spatial encoding by gradient fields and
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full signal detection.
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However, much shorter signal lifetimes can be observed for nuclei in molecules with more restricted mobility. Some tissues, including tendons, ligaments, bone, and teeth, are stringy or solid and owe their short T2 to low water mobility. 1H and other nuclei bound in large molecules such as lipids and
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proteins can be even less mobile and thus exhibit yet shorter T2. However, loss of macroscopic transverse magnetization can also stem from dephasing due to B0 inhomogeneity at the microscopic
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scale. A special case is lung tissue, in which rapid decay of the water signal associated with strong, microscopic variation in magnetic susceptibility is further influenced by diffusion. Generally, the exact mechanism of coherence loss is relevant when considering methodology. However, in the context of imaging, in particular in vivo, many of the main considerations hold equally for different
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types of signal decay. Therefore, unless stated otherwise, all kinds of coherence loss (such as T2 or T2*) will be jointly referred to as T2 decay in the scope of this work.
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Rapid signal decay hampers MRI in two ways. In conventional scans, signals with T2 below approximately one millisecond are largely missed during the dead time between MR excitation and detection, causing voids in resulting images. In addition, even if short-T2 signal is still captured, signal decay during data acquisition translates into apodization in k-space, entailing loss of spatial resolution
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or even artifacts.
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To expand the scope of application of MRI to short-T2 tissues, continuing efforts have been made to
image such components by means of dedicated techniques. In fact, the field of short-T2 imaging is
almost as old as MRI itself. In solid-state NMR, short T2 is often addressed by line-narrowing techniques [1-4]. Unfortunately, line-narrowing is barely applicable in vivo since it involves either rapid sample rotation or high levels of radiofrequency (RF) power absorption. Therefore, short-T2 imaging in vivo must primarily rely on sufficient speed of spatial encoding and data acquisition. The chief traditional strategies to this end are single-point imaging (SPI) [5] and ultra-short echo time (UTE) imaging [6], which both employ efficient center-out k-space encoding. Over the past decade, the field has been expanded by zero echo time (ZTE) methods [7], of which an array of
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implementations is now in use while more variants continue to emerge. Promising results of recent applications have promoted general interest in short-T2 imaging [8] and instrument manufacturers are increasingly following up with commercial implementations of new advances. At present, short-T2 MRI is a diverse field, and it can be a challenge to gain an overview and keep track of ongoing developments. Nevertheless, all of the techniques that are involved have emerged
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from a finite set of basic ideas that guide sequence design, signal detection, data processing, and
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hardware requirements. The goal of this article is to review the existing variety of short-T2 MRI methods from such a conceptual point of view, to work out commonalities and key differences
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between techniques, and to place them in a shared framework. It strives to offer guidance for getting started in short-T2 method development and for choosing the right approach towards a given
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application.
The presentation starts with a brief recapitulation of the physics underlying signal decay in MRI. It
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then discusses the basic principles of short-T2 imaging, which serve as a framework for a subsequent overview of existing techniques. Particular attention is given to two classes of techniques that are
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currently most employed in vivo, the UTE and ZTE approaches. A separate section is dedicated to specific hardware requirements. The final section of this review surveys the realm of current and potential applications.
2.1
Preamble
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2 Relaxation in a nutshell
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The general topic of relaxation is a complex branch of NMR physics. For the purpose of this article, no attempt is made to survey it fully or at significant depth. The aim of this section is rather to set the scene for MRI with short-lived signals by a brief, simplified review of the chief causes of rapid signal
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decay. For a thorough treatment of the topic we refer to NMR textbooks [9] as well as original work
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[10-12] and pertinent reviews [13-16]. 2.2
Overview
Relaxation is the process that drives a perturbed spin system back to thermal equilibrium with zero transverse
magnetization,
,
and
finite
longitudinal
magnetization,
.
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phenomenological, mono-exponential description of relaxation is given by the Bloch equations. In the absence of an RF field they reduce to
(1)
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(2)
with relaxation times T1 and T2, which vary greatly across different nuclei and their environments in given tissues and materials. The purpose of relaxation theory is to describe the mechanisms underlying relaxation and to relate T1 and T2 to more fundamental physical properties.
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In most practical cases, relaxation of nuclear spins is based on two key ingredients: 1) some magnetic
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interaction of the spins in question with their surroundings that depends on orientation or position of the molecules containing the spins and 2) quasi-random molecular tumbling. Tumbling motion
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translates interaction into magnetic field fluctuations at the spin locations. These fluctuations cause continuous slight variations in the precession frequency of the spins and in their angle with respect to
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the static field B0, leading to irreversible loss of transverse magnetization and recovery of longitudinal magnetization, respectively. Interactions
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2.3
Relevant interactions occur within molecules as well as between spins and their wider environment.
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To contribute to relaxation, interactions must be spatially anisotropic or inhomogeneous. In particular, intra-molecular interactions must depend on the orientation of the molecule with respect to the B0 field whereas inter-molecular interactions must change with distance or relative orientation. The
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usually strongest and therefore most important interactions are: 1) Dipolar coupling: The dipolar field associated with a spin’s magnetic moment contributes to the local field at the sites of nearby spins. Dipolar coupling is a through-space effect that drops rapidly
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with distance and therefore acts mostly intra-molecularly. It will be discussed in more detail below. 2) Chemical shift anisotropy: In electron shells of molecules, the external B0 field induces currents, which in turn add to the total field experienced by nuclear spins. This effect is generally anisotropic
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because the topology of a molecule favors certain electron paths. The strength and direction of the
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created field thus depend on the orientation of the molecule relative to B0. With respect to relaxation, this effect is most relevant in solids. In liquids and gases, it is typically canceled by motional averaging. 3) Electric quadrupole coupling: In nuclei with spin > 1/2, the distribution of charge is non-isotropic,
giving rise to an electric quadrupole moment that interacts with electric field gradients generated by surrounding electrons. This interaction depends on the relative orientation of the molecule and the nuclear moments and thus fluctuates upon tumbling in the B0 field. The effects of orientation-dependent interactions are schematically illustrated in Fig. 1 for the case of dipolar coupling between two equivalent nuclear spins in a molecule. Upon rotation of the molecule,
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the spins remain aligned with B0 such that their dipolar interaction changes (Fig. 1a). The associated effect on NMR spectra depends on the configuration and mobility of molecules in a sample (Fig. 1b). For a rigid single crystal, all local lattices form the same angle with B0, leading to a single sharp peak. A second symmetric peak is created by opposite spin alignment. In a powder, all lattice orientations co-exist and give rise to a continuum of resonance frequencies forming the characteristic Pake pattern
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of a powder spectrum [17, 18]. In liquids, tumbling of molecules partly cancels dipolar interaction by
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temporal averaging. In the limit of rapid tumbling, the spectrum converges to a single sharp peak as would be obtained without any coupling. However, in real liquids, tumbling occurs on a finite time
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scale, still averaging dipolar coupling and leading to a single spectral peak, yet with a finite linewidth (Fig. 1b), which is associated with relaxation as discussed in the following section.
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Free induction decay (FID) signals corresponding to these spectra are shown in Fig. 1c. The hypothetical sharp peak corresponds to a signal of constant amplitude. With relaxation, the signal decays exponentially. As this decay is driven by stochastic fields, it is irreversible and cannot be
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countered by spin-echo refocusing. The two sharp peaks of the single crystal correspond to oscillation with a constant envelope in the time domain while the broad, sharply delineated spectrum of the
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powder entails a rapid signal drop with subsequent ringing. In the ideal powder, the loss of coherence is not of stochastic nature. In principle, it could be fully recovered by refocusing and is thus not considered actual relaxation. Mechanism of relaxation
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2.4
The dependence of relaxation on random molecular motion is shown in the final plot of Fig. 2
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(“Relaxation”, bottom right): In liquids, high molecular mobility leads to long T2 and T1 whereas in solids restricted mobility causes short T2 associated with long T1. The whole Fig. 2 summarizes the mechanism of relaxation driven by orientation- or position-dependent interaction in combination with
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tumbling motion. Quasi-random rotational or translational motion and the anisotropic or non-uniform field associated with interaction lead to a random field ΔB(t) at the location of a considered spin. The
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amplitude of the field fluctuations is given by the interaction whereas the temporal behavior is linked with the speed of the tumbling motion and characterized by the correlation time τc. The latter can be
determined by calculating the autocorrelation function of the random field which is usually well described by an exponential function. The Fourier transform (FT) of the autocorrelation function is the spectral density J(ω) of the tumbling motion, i.e., the energy available as a function of frequency. It has the shape of a Lorentz curve and depends on τc as illustrated by two examples. The spectral density is the key link between molecular motion and relaxation. Relaxation changes the state of the spin system. Longitudinal relaxation affecting Mz is associated with change of the populations of the different energy levels, as indicated for a two-spin system in
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Fig. 2. The transition probabilities depend on the order of the state change and are proportional to the spectral density of the tumbling motion at multiples of the Larmor frequency ω0. 2.5
T1 and T2
The final plot in Fig. 2 shows the qualitative behavior of relaxation times as functions of the
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correlation time of tumbling motion. At very short τc as found in small-molecule liquids, interactions are averaged effectively, resulting in very slow relaxation, i.e., long T1 and T2. As τc increases, both
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types of relaxation initially speed up. T1 reaches a minimum at τc ≈ 1/ω0, reflecting dependence on
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J(ω0) and J(2ω0). Hence, longitudinal relaxation is most effective when 1/τc approximately matches the Larmor frequency, similar to the resonance condition for RF excitation. It is instructive to view
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longitudinal relaxation as caused by the transverse components of the random field whose action on Mz starts to cancel itself once spins perform more than half a precession cycle during the correlation time [14]. In contrast, T2 has no minimum, reflecting the fact that for large τc it is dominated by J(0).
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In this regime, transverse relaxation is driven mostly by the longitudinal component of the random field. Longitudinal relaxation inherently entails a certain loss also of transverse coherence, limiting
2.6
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the transverse relaxation time to T2 ≤ 2T1. In most practical cases T2 ≤ T1 [19]. T2*
In real experiments, the decay of an FID signal obtained from a given volume is often faster than
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according to T2. The chief cause is spatial variation of the static magnetic within the volume. The associated frequency spread causes dephasing of the spins, i.e., loss of macroscopic transverse coherence and thus signal decay. As long as the field variations are static, the dephasing can, in
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principle, be undone by spin-echo refocusing. In contrast to pure random-field-driven relaxation, signal decay including macroscopic dephasing has been termed apparent transverse relaxation. The shape of such decay varies widely depending on the underlying field distribution. Nevertheless, it is
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frequently approximated by an exponential with the reduced time constant T2*. In the following, three
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cases of apparent transverse relaxation are discussed more closely: 1) Dipolar coupling in a powder as described in Fig. 1. In this case, the shape of the FID deviates strongly from exponential behavior. Large frequency spreads on the order of typically tens to hundreds of kHz require large-bandwidth, high-power RF pulses for refocusing. Alternatively, the dipolar interaction may be averaged out by spinning the sample at the magic angle or eliminated by RF decoupling pulses [18]. Note that such approaches are usually impractical in vivo. 2) Microscopic structure: Biological tissues (e.g. trabecular bone) and other samples of interest (e.g. porous rock) often exhibit microscopic structure on a length scale below feasible MRI
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resolution. Related microscopic variation of magnetic susceptibility causes corresponding field inhomogeneity and signal dephasing. In T2*-weighted MRI, the associated signal decay alters image intensity on a per-pixel basis and can be used to characterize the underlying structure and materials involved. 3) Macroscopic B0 inhomogeneity: Macroscopic compartments filled with materials of differing
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bulk susceptibility such as soft tissue, bone, and air, give rise to long-range field variation
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that also causes dephasing and signal decay. However, such dephasing is not characteristic of the tissue in which it occurs. For compartment sizes in the order of or larger than the spatial
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resolution, macroscopic susceptibility effects rather hamper MRI by depiction errors, e.g., signal shifts or blurring, and unspecific signal dropout. These problems can be mitigated by
2.7
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increasing spatial resolution and the strength of gradient fields used for spatial encoding. Diffusion
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In the presence of static microscopic field inhomogeneity, signal decay is further influenced by diffusion of the molecules that bear the resonant spins. Random motion translates field inhomogeneity
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into random field fluctuation that is individual for each spin and thus affects the nature of dephasing among the spin ensemble. This mechanism differs from relaxation as discussed above in that the effective field fluctuations are smaller by many orders of magnitude and arise from fields on a greater length scale such that their correlation times are much longer. As a result, dephasing under diffusion
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is amenable, at least in part, to refocusing by radiofrequency pulses. In a linear field gradient with amplitude G and unbounded Gaussian diffusion, the signal amplitude of an FID decays according to
(3)
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with self-diffusion constant D and gyromagnetic ratio γ [20, 21]. In spin-echo experiments with echo
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time TE, refocusing partly recovers the decay, resulting in an apparent relaxation rate of
(4)
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Hence, MR experiments may be hampered by increased apparent relaxation due to diffusion but can also give access to the physical property D. 2.8
Examples
As discussed above, various mechanisms can cause loss of coherence and thus rapid signal decay in tissues and materials. Accordingly, short-T2 MRI has a large variety of potential imaging targets, including the following prominent examples that will be further discussed in Section 10:
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1) Solids: In solids, molecular motion is strongly restricted, leading to long correlations times. As seen in the relaxation plot of Fig. 2, in such substances short T2 is associated with long T1. The latter has two key implications for sequences with rapidly repeated RF excitations as introduced later: 1) steady-state magnetization is generally low and 2) maximum steady-state signal is reached with small excitation flip angles. In the biomedical context, a solid of
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particular interest is hydroxyapatite in the bone matrix [22].
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2) Bound water: Water molecules are small, leading to short correlation times and slow relaxation in free water. However, when bound to macromolecules or solids, water molecules
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exhibit reduced mobility and much shorter T2. An example is collagen-bound water in bone [23].
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3) Water in collagen fibers: Some tissues (e.g. tendons, cartilage) are composed of collagen fibers which are organized in directed, parallel bundles. Motion of water embedded in such highly ordered structures is strongly restricted, leading to short T2. In addition, there are
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preferred orientations along which bound water molecules line up. Thus the effect of dipolar coupling and hence T2 become a function of the angle between fibers and B0 [24-33]. In T2-
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weighted acquisitions, signal varies and peaks at the so-called magic angle of 55°. This needs to be considered for image interpretation but also holds potential for contrast manipulation. 4) Myelin: This membrane tissue, which plays a key role in the nervous system (see section 10), is largely composed of phospholipids arranged in liquid crystalline bilayers. Intra-molecular
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dipolar interactions cause rapid T2 decay of 1H signal, which varies with the orientation of the bilayer with respect to B0. As in typically probed volumes all orientations exist, the resulting
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signal is well described by a superposition of exponential decays with a wide range of T2 values, leading to so-called “Super-Lorentzian” lineshapes [34, 35].
5) Lung parenchyma: Signal decay of water in pulmonary tissue is an important example for the diffusion-driven mechanism described in section 2.7. Microscopic tissue-air interfaces at the
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level of the alveoli create local field gradients that rapidly dephase spins in water diffusing in
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the tissue [36]. This results in fast signal decay in gradient-echo techniques, which can be partly reversed with spin-echo sequences, according to Eqs. 3 and 4.
3 Fundamentals of short-T2 MRI
In this Section, the specific requirements for depicting tissues with short T2 are expressed in terms of two guiding principles. Starting from these principles, concepts of sequence design are derived, leading to four basic templates of short-T2 techniques.
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3.1
Principles
Figure 3 illustrates the challenges of short-T2 imaging. Common echo sequences hardly capture any signal from short-T2 spins due to signal loss by a factor
until the echo is formed and acquired
around the echo time TE (Fig. 3a). For gradient echoes this holds equally for pure T2 decay and T2*
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decay. In spin echoes, the latter is partly refocused, which can be beneficial when T2* is much shorter than T2. However, at very short T2*, RF refocusing requires commensurately large bandwidth and thus
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power deposition, which limits the applicability of spin echoes for this purpose. Therefore this option
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is not further considered here in favor of techniques that can equally handle short T2* and T2 . To contain the signal loss observed with echo techniques, dedicated short-T2 sequences must start
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gradient encoding and data acquisition quickly after signal excitation. Following this first principle renders short-T2 materials visible to begin with (Fig. 3b). However, signal decay during data acquisition still impairs spatial resolution. In mathematical terms, this is due to apodization in k-space,
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which is equivalent to convolution with a kernel of reciprocal width in the image domain [37]. To avoid such resolution loss, all image data must be acquired within a small time range after excitation.
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In Fig. 3c, this second principle is implemented by use of a stronger gradient, which shortens the readout and thus improves the sharpness of short-T2 features. More formally, the relationship between time and k-space in an MRI sequence is described by
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defining the time of acquisition as the time after signal excitation at which a k-space point is acquired. In echo-based MRI techniques, the value of this function at the k-space center, , is the echo time TE. The same definition of TE has widely been adopted for short-T2 imaging even though
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most of the techniques work without actual echo formation (see Section Error! Reference source not found. for details). To keep with this convention, the same notion is used in the present work. The range of times across which data are collected shall be referred to as the acquisition range
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!" .
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Based on these definitions, the principles of short-T2 sequences can be narrowed down to 1) data acquisition must start early, i.e., !" must be small, and 2) the acquisition range must be
small. Both of these conditions must hold on the time scale of relevant T2. Table 1 summarizes these considerations along with characteristic sequence parameters and their impact on image quality. Note that !" in short-T2 imaging is often, but not necessarily, equal to TE. To approach the encoded, nominal resolution, the acquisition range should approximately match the decay time constant, i.e., # [38]. Shortening the acquisition further will yield diminishing returns in terms of resolution but comes at the expense of signal-to-noise ratio (SNR).
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The principles have an immediate consequence on any short-T2 sequence, as illustrated in Fig. 4. Competing with fast signal decay by reducing TE and/or requires rapid signal encoding. On the other hand, achieving high spatial resolution requires large time integrals of gradient amplitude. These demands can be fulfilled jointly only by using strong gradients and corresponding large spread of resonance frequency across the imaged object. As discussed in the following, such strong gradients
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may be active during RF excitation and/or data acquisition, making large bandwidth a characteristic of
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short-T2 MRI. This consideration also reveals the fundamental challenge that necessarily short observation times of rapidly decaying transverse magnetization are equivalent to high-bandwidth
Concepts
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3.2
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acquisition. Hence, short-T2 imaging is generally more SNR-limited than conventional MRI.
At the level of sequence design, the principles discussed above lead to three chief concepts that recur across the variety of short-T2 techniques:
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A) No slice selection: Slice selection under a gradient of finite strength requires RF pulses of suitably small bandwidth and thus long duration, followed by slice rephasing. To avoid the associated delays,
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short-T2 imaging frequently works with non-selective volume excitation, omitting the slice selection gradient. All spatial encoding is then achieved by three-dimensional (3D) Fourier encoding. (An exception to this approach will be described later in Section 4.3.)
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B) Radial center-out encoding: To reach every targeted k-space point in the shortest possible time, kspace is traversed by straight radial trajectories that point outward from the center. This approach is
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used in two principally different ways:
Pure phase encoding: after each excitation, a single point in k-space is targeted and sampled once reached.
Pure frequency encoding: data acquisition is performed continuously along radial trajectories.
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C) Excitation under encoding gradient: To avoid losing time ramping up the readout gradient after
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excitation, the gradient can be switched on first such that it is already present at when excitation is performed. Figure 5 visualizes the combinations of these design elements and indicates published implementations using the respective approaches. While all techniques share concepts A and B, they differ with respect to the options for concepts B and C. All current short-T2 techniques can be represented in this scheme and thus viewed in a shared conceptual framework. As the figure indicates, two options each with respect to concepts B and C give rise to four basic sequence templates, which will now be considered in more detail.
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3.3
Basic sequences
The four basic schemes of short-T2 sequences and their associated k-space sampling patterns are shown in Figs. 6 and 7, respectively. All techniques employ non-selective hard-pulse excitation followed by 3D radial center-out encoding in k-space. Pure phase encoding is used in constant time imaging (CTI) [39] (Fig. 6a) and SPI [5, 40, 41] (Fig.
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6b). In both, a gradient of varying amplitude and direction is used to travel to one point in k-space per
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excitation (Fig. 7a). Each k-space point is sampled at the same fixed = TE, which must be chosen large enough to permit reaching the maximum k-space excursion that corresponds to the targeted
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resolution. While k-space is traversed radially, the points sampled are typically chosen to lie on a
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Cartesian grid. By virtue of fixed , all data points have identical T2 weighting. Therefore, signal decay prior to TE does not impair spatial resolution. However, it does affect the SNR in resulting images, most so for short-T2 material. In CTI, TE includes time for ramping the encoding gradient up
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and down such that it is off during excitation and acquisition. This has the advantage that the bandwidth of both can then be limited to the intrinsic frequency spread in the object. However, it
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exacerbates the SNR penalty of signal decay.
As a remedy, in SPI TE is reduced by performing RF excitation and data acquisition in the presence of the gradient (Fig. 6b). The bandwidth of both must then be increased by making them suitably short. As a result, the SNR yield of SPI is subject to competing influences of signal decay and acquisition
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bandwidth. It is optimal at , requiring a maximum gradient amplitude of $ %&
for given T2 and spatial resolution %& [40]. In addition to its short-T2 capabilities, the SPI approach
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offers relatively silent operation if gradients are not switched off between excitations but merely adjusted to a nearby value [42]. The key advantage of purely phase-encoded techniques is full nominal spatial resolution at short T2.
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Their chief drawbacks are long scan times and low SNR efficiency associated with point-wise sampling, low acquisition duty cycle, and comparatively long TE. For these reasons, pure phase
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encoding is rarely used in vivo [43]. It is mainly applied for imaging materials [44, 45] and, in MR microscopy, to mitigate diffusion effects [46]. Recently, as discussed in Section 6, partial SPI acquisition has found additional use as a means of complementing other, more time-efficient techniques. The terminology of purely phase-encoded imaging is not entirely consistent, with somewhat variable use of the terms CTI and SPI in the literature. With pure frequency encoding, lines in k-space are sampled by continuous signal acquisition during radial encoding. In this fashion, shorter TE, faster overall scanning, and higher SNR efficiency are accomplished. For an image matrix size M, the gain in acquisition duty cycle relative to SPI leads to
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an increase in SNR efficiency of at equal maximum G and of at optimized G (see 9.2). In both UTE [6, 47-54] (Fig. 6c) and ZTE [7, 55-58] (Fig. 6d) imaging, radial encoding is performed by playing out gradients of fixed amplitude but direction varying from repetition to repetition. Thus one spoke is obtained per excitation (Figs. 7b and 7c). In the UTE sequence, data acquisition is started concurrently with ramping up the gradient shortly after excitation. The minimum TE then amounts to
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half the RF pulse duration plus the dead time required to change from transmit to receive operation.
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Due to gradient ramping, the central k-space is traversed at lower, linearly increasing speed (Fig. 7b). This is avoided in the ZTE sequence where RF excitation is performed only after switching on the
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gradient, thus achieving instant, full encoding speed. As a consequence, similar to SPI, RF excitation must be broadband and near-silent operation is possible by only adjusting and not switching off the
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gradient between excitations. In the ZTE case, spatial encoding starts immediately at the time of creating transverse magnetization; hence the time TE at which it leaves the k-space center is effectively zero. However, during the finite RF pulse duration and the RF dead time initial signal is
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lost, leading to a gap in central k-space (Fig. 7c), which needs to be addressed by additional measures (see 5.3 and 6.1).
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For in-vivo MRI, UTE and ZTE imaging and their relatives are currently the most suitable and most widely used techniques and will therefore be discussed in more detail in Sections 4 and 5. For all implementations discussed above, the concept of radial encoding is central to short-T2
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depiction performance. With common Cartesian encoding, the principles stated initially cannot be implemented as effectively. Notwithstanding this, Cartesian gradient echo imaging methods can also
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be optimized for short-T2 imaging [59, 60]. 4 Ultra-short echo time imaging The UTE strategy is the most established and most widely deployed short-T2 approach in use today.
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While arguably simple in principle, UTE sequence design and image reconstruction deviate significantly from standard MRI. To obtain robust results, implementations must consider a number of
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subtleties and pitfalls as detailed below. 4.1
Gradient ramping
The basic idea of the UTE sequence as shown in Fig. 6c is to start data acquisition very soon after RF excitation while concurrently ramping up the radial readout gradient to its target amplitude. The gradient ramp plays a key role in UTE MRI because it affects image quality in several ways. Depending on the desired gradient strength and the available slew rate (see Section 9.2), the ramping part can engage a considerable portion of the total encoding time. For a given nominal spatial resolution, the encoding time must be extended by half the ramp duration relative to the ZTE case
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with full gradient amplitude throughout. The effect of the gradient shape on depiction behavior in UTE imaging is illustrated in Fig. 8, using the point spread function (PSF) as a criterion. The time of acquisition of each k-space value (Fig. 8a) shows sections of increasing and constant k-space speed during gradient ramp and plateau, respectively. Longer ramps lead to larger acquisition range . In conjunction with T2 decay, the acquisition timing translates into the modulation transfer function
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(MTF) (Fig. 8b), which is then Fourier-transformed to obtain the PSF (Fig. 8c). As ramp times
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increases, apodization of the MTF causes progressive broadening of the main lobe of the PSF. The
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normalized display (Fig. 8d) illustrates the associated loss of spatial resolution. Gradient fidelity
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Gradient ramping is also critical to UTE image quality in that the demanded gradient waveforms may not be produced with the necessary fidelity. Delays, finite gradient amplifier bandwidth, and eddy currents cause deviations from nominal gradient dynamics. These may include dynamic fields
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different from first spatial order, usually with a dominant zeroth-order component. The simulations in Fig. 9 illustrate effects of eddy currents on UTE image quality. The distorted gradient shape in Fig. 9c
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leads to severe alterations in signal intensity when image reconstruction is based on the nominal trajectory (Fig. 9d). So-called B0 eddy currents that generate zeroth-order field and thus spatially constant phase errors (Fig. 9e) result in edge artifacts (Fig. 9f). Both types of error are amenable to retrospective correction. Phase perturbation by B0 eddy currents can be removed from the acquired
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raw data while first-order field effects can be addressed by image reconstruction based on the effective, distorted k-space trajectory. In this way, a virtually artifact-free image is obtained,
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comparable to imaging with ideal gradient behavior (Fig. 9b). Models of system imperfection for retrospective correction may be set up at different levels of complexity ranging from mere gradient chain delays to general first-order field perturbation,
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additional zeroth-order [61], and second- and higher-order field terms [62]. Various strategies have been described for determining the parameters of such models, including approaches based on raw
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UTE data itself [63, 64], calibration scans [65-67], impulse-response characterization of the gradient system [68], and measurement of field evolution using modified MRI sequences [69, 70] or dedicated field probes [71] (see Section 9.5). 4.3
RF excitation
In the 3D UTE sequence shown in Fig. 6c, spatially non-selective RF excitation is followed by 3D radial encoding of the whole object. Alternatively, unlike SPI and ZTE techniques, UTE permits slice excitation under a separate selection gradient, forming a two-dimensional variant. To avoid increasing TE by longer RF pulses and slice refocusing, it has been proposed to employ half-excitation pulses
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with opposite slice gradients in two successive acquisitions [47, 72, 73]. Full slice selectivity is then obtained by summation of acquired data. On the downside, besides doubling the scan time this approach has been reported to increase sensitivity to eddy currents [74, 75] and motion. Relaxation during excitation must be considered when pulse durations are significant relative to T2 of interest. In-pulse relaxation reduces the actual flip angle and renders it T2-dependent [52]. The
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presence of an RF field changes the physics of relaxation, which must be taken into account in
Echo time
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4.4
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sequence optimization and the interpretation of data [76, 77].
Two different definitions of TE are common in the UTE literature. The first, as shown in Fig. 6c and
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used throughout this article, assumes that relaxation starts at the magnetic center or focus point of the RF pulse [77, 78]. For symmetric waveforms, including the commonly employed hard pulses, this is the center of the pulse. The second definition assumes that relaxation starts only after the RF pulse
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[38]. Both pictures are incomplete in that they do not account for relaxation during excitation as discussed in the preceding section. Hence, the full characterization of a short-T2 sequence also
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requires RF shape and duration in addition to TE of either definition.
In short-T2 samples, the SNR that can be achieved with UTE imaging depends on the shortest possible TE, which is governed by sequence implementation and hardware constraints (Tab. 1). A lower bound
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for TE is given by the RF dead time, which is composed of the fraction of the RF pulse duration after its magnetic center, the time required to switch between RF transmission and reception, and the leadin time of bandpass filtering [79] (see 9.3). On current MR systems, typical TE values for 3D UTE
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MRI are in the order of several tens of microseconds. In the UTE scheme, TE can also be varied, which is a useful feature. It permits creating T2* contrast, emphasizing short-T2 tissues by image subtraction [80], and mapping T2* quantitatively [81].
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Deliberate choice of somewhat longer than minimal TE can serve to avoid detection of signals that
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decay too fast for meaningful spatial encoding even with the UTE approach. Disturbing signal of such exceedingly short T2 most commonly originates from hardware parts and causes cloudy image background when unaddressed [82] (see 9.3). 4.5
Image reconstruction
Due to non-Cartesian sampling, image reconstruction from UTE data involves interpolation in kspace, which is usually performed with gridding procedures [83, 84]. In particular, this includes correction for non-uniform k-space density which can be accomplished by using a pre-calculated density correction function (DCF) [85-89]. One approach to obtain a DCF defines local k-space
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density as the area surrounding a sampled point as calculated by Voronoi diagrams [86]. For radial sampling, this approach leads to the so-called rho filter [90] for which an implementation is provided in Section 12.1. Density correction is an approximation based on splitting up reconstruction into FT and a preceding filter, the DCF. More rigorously, reconstruction is viewed as a general inverse problem. It can be
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solved efficiently using iterative algorithms [91] which inherently perform correction for non-uniform
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k-space sampling. In this case, application of a DCF can serve for pre-conditioning, thus accelerating convergence. The choice of the DCF then affects only convergence but not the accuracy of the
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reconstruction result.
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5 Zero echo time imaging
Image acquisition according to the ZTE scheme shown in Fig. 6d dates back several decades. Renewed recent attention has been driven by technical and methodological advances as well as
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increasing interest in short-T2 capability and silent MRI. To express its relation to the more established
5.1
Sequence properties
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UTE method, the approach has been dubbed ZTE imaging [7].
The basic ZTE sequence is an imaging analogue of the pulse-acquire NMR experiment. The only modification consists in spatial spreading of resonance frequencies by addition of a projection
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gradient that is present during RF excitation and FID acquisition. This approach was in fact used in the very first MRI experiment ever reported [92]. Only later it was identified and exploited as an
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approach particularly suitable for short-T2 imaging [55-58]. Relative to UTE and standard present-day MRI the distinguishing feature of ZTE imaging is RF excitation in the presence of the encoding gradient. While arguably a small difference in a sequence
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diagram, it is of great consequence with regard to sequence considerations, hardware requirements, data acquisition, and image reconstruction. The chief objective of excitation under the encoding
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gradient is to remove the need for gradient ramping while short-T2 magnetization evolves. It saves
time, offers immediate full k-space speed, and minimizes gradient switching even after data acquisition. In this way, enhanced short-T2 sensitivity is obtained in combination with reduced
acoustic noise or even virtually silent operation and robustness against eddy currents. Removal of gradient ramps also reduces the time required per repetition and increases the acquisition duty cycle. In combination with large gradient strengths as required for short T2, very short repetition time (TR) can be achieved and scan times for ZTE imaging are comparatively short [93].
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These capabilities come at the expense of shifted and partly increased hardware requirements. Continuous gradient operation for short-T2 imaging requires high gradient strength at full duty cycle. This demand is not commonly met by present-day gradient systems and has prompted developments in dedicated gradient hardware (see Section 9.2). On the other hand, the ZTE mode comes with reduced specifications in terms of gradient switching speed. In turn, it partly shifts hardware
Dead time
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5.2
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transmit-receive (T/R) switching are essential as will be detailed in the following.
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challenges from the gradients to the RF subsystem. In particular, high-bandwidth excitation and rapid
In pulsed MR experiments, signal excitation and acquisition are separated in time to accommodate the
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large difference in the RF amplitudes involved. Changing between the respective modes of operation takes time, resulting in a dead time after RF excitation during which MR signal is not yet detected. In many MR experiments, especially those involving echo formation, acquiring data immediately after
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RF excitation is not of interest. However, in short-T2 imaging, encoding and acquiring fresh MR signal as soon as possible is one of the key principles (see 3.1). With immediate encoding at full k-
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space speed, the ZTE sequence is arguably the most rigorous implementation of this rule and thus particularly vulnerable to data loss during the dead time. In 3D k-space it causes a spherical gap at the center (Fig. 7c). To contain the gap size one must minimize the dead time, which has contributions from the RF pulse, T/R switching, and digital filtering (Fig. 10) [7]. The chief means of limiting the
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gap are thus the use of short RF pulses [93] (see Section 5.5), speeding up T/R switching (see Section 9.3), and the design and use of short digital filters (see Section 9.4). The residual k-space gap must be
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addressed either by means of image reconstruction (see Section 5.3), by acquisition of additional data (see Section 6.1), or a combination of the two. Ideally, RF dead time would be avoided altogether by performing signal acquisition also during
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excitation. However, due to the large dynamic range between transmit and receive signal this
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approach is of limited practicality (see Section 9.3). 5.3
Algebraic Reconstruction
Straightforward image reconstruction from data with a gap in central k-space, using standard FT, leads to prohibitive artifacts of low spatial frequency. However, for small gap sizes the missing data can be recovered by taking two measures: 1) oversampling in the direct, radial dimension (Fig. 7c, Fig. 10) and 2) treating reconstruction as a general inverse problem set in the framework of linear algebra [94]. Standard sampling according to the Nyquist-Shannon theorem is associated with harmonic encoding functions that are mutually orthogonal within the field-of-view (FOV). In contrast, oversampling is associated with encoding functions that exhibit non-vanishing scalar products with
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the Nyquist set. In the algebraic framework, these dependencies permit the implicit formation of missing low-frequency harmonics by linear combination of available encodings [95]. This approach is known as finite-support or band-limited extrapolation and was previously used for low-frequency restoration in MRI in a different context [96]. It is limited to a relatively small range of extrapolation, beyond which it becomes impractical due to ill-conditioning (see Section 5.4). Greater oversampling
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improves conditioning somewhat but also increases the amount of resulting data. Therefore,
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oversampling factors range between 2 and 4 in practice [95].
Formally, the algebraic approach models data acquisition by means of an encoding matrix [97]. Its
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rows reflect the encoding functions actually implemented. Hence, the gap is represented by the absence of corresponding rows while oversampling elsewhere leads to an increased number of rows
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relative to standard Fourier encoding [7]. Finite support is implied by limiting the length of the rows to the FOV. Algebraic reconstruction is achieved by solving the encoding equation for the unknown image, eliminating the dead time gap and returning to regular sampling density in k-space. Figure 11
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illustrates this approach for the case of one-dimensional (1D) ZTE data. In principle, it can be straightforwardly expanded to 3D, working with one single encoding matrix. However, the latter will
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usually be too large for direct inversion and require iterative treatment. Alternatively, and more efficiently, algebraic processing can be restricted to 1D sub-problems, reconstructing full k-space lines through the origin from pairs of acquisitions with opposite gradient direction. The resulting full
3D image [7].
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set of radial k-space data is then merged by standard 3D gridding (see Section 4.5) to form the final
Data extrapolated to the gap in the described manner reflect the actual T2 decay during the dead time.
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In particular, the value at k = 0 is not affected by transverse relaxation. Hence, although these data points are not actually acquired, they correspond to signal obtained as if sampling was started at TE =
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0, which has led to the “ZTE” notation. 5.4
Bandwidth limitations
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There are limitations to the gap size and hence the signal bandwidth that can be handled with the algebraic approach. These arise from deteriorating conditioning of the encoding matrix as gap size increases [94]. Ill-conditioning boosts error in the signal and the signal model, causing three kinds of image distortion: 1) enhancement and spatial correlation of noise, 2) aliased and amplified out-ofband signal, and 3) intra-voxel oscillation of the spatial response [95]. Usually, the first two are dominant in practice and will therefore be described in more detail below. As the reason for illconditioning is data missing in central k-space, all image distortions described are of low-frequency nature and may not be distinguished easily.
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1) Noise in common Cartesian k-space raw data exhibits constant variance which leads to the usual uncorrelated appearance of noise in MR images. In contrast, in reconstructed ZTE data, noise variance increases strongly towards the k-space center, thus causing noise correlation in image space. Therefore, at large gaps, noise in ZTE images can have an unusual artifact-like appearance (Fig. 12). Its random nature is only observed by varying shapes in the image resulting from different noise input
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data. Importantly, ZTE image reconstruction benefits from intrinsic averaging in central k-space
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associated with the sampling density of 3D radial data increasing strongly at small radii. Therefore, distortions appearing in 1D projections do not necessarily translate into deterioration of 3D image
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quality.
2) Out-of-band signal violates the fundamental assumption of ZTE reconstruction in that it stems from
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locations outside the FOV. As a consequence, such signal will be assigned into the FOV, corresponding to aliasing. Furthermore, related to the bad conditioning it will appear strongly amplified in the image. This is a serious practical issue as ZTE acquisition not only captures the
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desired short-T2 signal of the object of interest but also unwanted contributions from nearby hardware parts, in particular RF coils (Fig. 13). The straightforward solution of targeting a larger FOV
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including all detected signal is of limited scope as it results in larger bandwidth, i.e., larger gap size, as well as longer scan times and larger data sets. Better ways of reducing such background signal are suppression techniques [7, 98], image subtraction [99], and designing RF coils with negligible content
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of the observed nucleus (see Section 9.3).
Within these boundaries and depending on setup and sequence parameters, practical gap sizes in ZTE experiments are usually limited to approximately 2 - 3.5 DW, where DW = 1/BW is the Nyquist dwell
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time for the bandwidth (BW) in the FOV. Using ZTE-optimized RF hardware, this can nevertheless translate into BW as high as 500 kHz [93], which is the key to high-resolution short-T2 MRI (see
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Section 3.1). 5.5
RF excitation
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RF excitation in ZTE imaging plays a particular role as it occurs while the radial encoding gradient is already on (Fig. 6d). To ensure consistency between data of different radial directions, the spectrum of the excitation pulse must be largely uniform within the bandwidth spanned by the gradient across the object. In the basic ZTE sequence, a short block-shaped, so-called hard pulse is used. For greatest uniformity the pulse duration should not exceed approximately DW/3 [100]. However, in practice durations up to about DW can also provide useful results. For bandwidths of hundreds of kilohertz this limits hard pulse durations to a few microseconds, which is much shorter than RF pulses commonly applied on
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clinical MR scanners. With the maximum B1 amplitudes usually available on such systems, flip angles of only a few degrees can be obtained (see Section 9.3). On the other hand, as little time is spent on RF pulses and gradient switching, the ZTE sequence can be operated with very short TR (see Section 7.1), even below 1 ms [93], thus enabling maximum steady-state magnetization even at very small flip angles.
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If flip angles should nevertheless be increased due to low B1 or very high bandwidth, or to create T1
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contrast, three options have been proposed, namely 1) longer hard pulses with intensity correction, 2) short modulated pulses, or 3) long modulated pulses with simultaneous excitation and acquisition.
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1) By increasing the hard pulse duration, its bandwidth will be reduced. For durations above
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DW, the main lobe of the Sinc-shaped pulse spectrum exhibits relevant variation in the FOV. Hence, larger flip angles are primarily obtained in the central FOV, which may still be useful if this coincides with the location of interest. Inconsistencies in signal intensity between
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projections can be reduced by correcting for the pulse profile [101]. This approach is limited by singularities above durations of 2DW and by variations of local steady-state magnetization
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over time.
2) The hard pulse is replaced by an RF excitation with both modulated amplitude and frequency [102, 103]. If designed appropriately, such pulses enable high spectral uniformity at high flip angle efficiency and thus outperform hard pulses of equal duration (Fig. 14). Unlike hard
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pulses, their length is not limited by uniformity considerations but by the maximum acceptable dead time (see Section 5.2). 3) Further extending modulated pulses may lead to intolerable dead times. To reconcile long
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pulses and short dead time it has been suggested in the SWIFT (sweep imaging with Fourier transformation) technique to overlap RF excitation and acquisition [104, 105]. This is indeed possible by alternating the two RF operation modes rapidly (Fig. 15), thus offering enhanced
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degrees of freedom in pulse definition. However, frequent switching comes at the expense of
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reduced acquisition duty cycle and thus SNR efficiency. Furthermore, feasible bandwidth is ultimately limited by the inverse of the dead time [79]. To image large bandwidths at reduced switching rate, the SWIFT approach has been expanded to a multi-band mode [106]. Alternatively, technically challenging approaches have been proposed where excitation and acquisition are performed truly simultaneously without gapping (see Section 9.3).
Apart from short hard pulses, all excitation options listed above generate some kind of spectral nonuniformity which needs to be considered during image reconstruction [101, 103, 104, 107]. A potential implementation obstacle may be the high temporal resolution required for creating highbandwidth pulses which is not always available on clinical scanners (see Section 9.4). A more
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fundamental limitation is posed by the large power deposited by high-bandwidth excitations, leading to potentially excessive specific absorption rates (SAR). 5.6
Comparison with UTE
In summary, the ZTE option exhibits the following advantages (+) and disadvantages (-) with respect
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to UTE imaging: Higher short-T2 sensitivity and resolution
+
Greater resilience against off-resonance effects
+
Greater robustness against eddy currents, no trajectory calibration required
+
Silent operation
−
High-bandwidth RF excitation: higher SAR and tighter limits on flip angles
−
Bandwidth limited by dead time
−
Only 3D
−
Limited intrinsic contrast
−
Higher sensitivity to background signal
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6 Intermediate techniques
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+
Various modifications of the basic sequences discussed so far in Sections 3.3, 4, and 5 have been proposed to mitigate some of their limitations. This results in the intermediate techniques indicated in
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the classification scheme of Fig. 5. They implement transitions on the two axes corresponding to concept B (vertical axis) concerning phase and frequency encoding and concept C (horizontal axis)
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concerning the succession of RF excitation and gradient switching. Intermediate techniques are obtained by either combining data obtained with different techniques or by merging opposed concepts within a sequence.
Filling the ZTE dead-time gap
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6.1
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Gap size limitations in algebraic ZTE have mainly been tackled by acquisition of additional data in the k-space center: 1) Sampling at k = 0 by additional acquisition in the absence of a gradient improves the conditioning [95, 108, 109]. However, this data point will have a finite TE equaling the dead time, thus leading to considerable signal decay for tissues and materials very short T2 values. This introduces an inconsistency with the bulk data which leads to image artifacts for badly conditioned algebraic reconstruction. 2) The previous approach is expanded in the PETRA (pointwise encoding time reduction with radial acquisition) technique [110] by filling the entire spherical void in k-space with
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Cartesian SPI data (Fig. 16a). In this way, arbitrary gaps can be handled, though at the expense of scan time and SNR efficiency. The effect on these two parameters increases steeply with gap size. 3) Alternatively, in the technique known as “water- and fat-suppressed proton projection imaging” (WASPI) [111], the dead-time gap is filled with radial data acquired at a gradient
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strength reduced by a factor 2 × dead time / DW (Fig. 16b), thus achieving Nyquist sampling
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in the very center also. By adapting the angular density to the radius given by the gap size, the additional scan time is usually less than for PETRA.
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The latter two options change the weighting obtained in k-space associated with T2 decay. The related effect of on image quality can be assessed via the PSF (Fig. 17). The PSFs show that compared to
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ZTE, the main lobe width is reduced for both options but the PETRA technique offers relatively benign side-lobe behavior whereas WASPI is prone to increased side lobes. It has been found that this tendency can lead to artifacts in the presence of components with very short T2 on the order of the gap
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size, stemming, e.g., from hardware parts (see Section 9.3) or solid tissues [109].
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4) In the hybrid filling (HYFI) technique [112], radial data from a WASPI acquisition with multiple gradient strengths and a Cartesian SPI acquisition in the very center similar to PETRA are combined. This approach enables improved PSF quality as compared to WASPI with less scan time increase than required for PETRA.
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In the techniques described above, data missing in k-space is complemented by performing additional gradient encoding. In a different approach following the principles of parallel imaging [97],
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supplementary information is obtained by detecting the MR signal with a receiver coil array. During image reconstruction, independent data from multiple coils facilitates deriving missing from existing neighboring data, thus improving the conditioning associated with large gap size [113]. A challenge of this approach is the need to solve the full 3D inverse problem for image reconstruction, differently
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from the simplified procedure outlined in Section 5.3. Finally, the encoding potential of RF pulses
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may be exploited to reduce the effective gap size [114]. 6.2
Reducing the excitation bandwidth
Another class of intermediate techniques departs from single gradient ramps either before or after the RF pulse by performing excitation under a lower gradient which is then ramped to full strength for rapid encoding. Such techniques have been proposed as gradient-modulated PETRA [115] or RHE (ramped hybrid encoding) [116] as well as gradient-modulated SWIFT [117]. As an advantage, RF excitation can be performed at lower bandwidth, thus relaxing flip-angle and SAR constraints. On the
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other hand, gradient ramping per se and its fidelity affect image quality in analogy to the UTE case (see Sections 4.1 and 4.2). 7 Further sequence aspects 7.1
Repetition time
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The sequences that have been described above are usually operated with short TR and a steady state of
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the magnetization. In this way, the many repetitions required for 3D acquisition can be performed within useful scan times. Assuming complete spoiling of transverse magnetization after each TR,
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maximal transverse steady-state magnetization ' and thus optimal SNR are obtained by excitation
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at the Ernst angle ( [118]. For ) * , which holds for most short-T2 imaging scenarios, both ( and ' are approximately proportional to ) (see Section 12.2). ' + ) implies that, when operating in the Ernst-angle regime, changing TR has no effect on SNR efficiency [119]. For example,
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reducing TR by a factor of two while doubling the number of averages will yield the same SNR at equal scan time. Furthermore, ( + ) implies that changing TR does not affect the SAR generated
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by the excitation.
To achieve short TR, gradient spoiling must be implemented in an efficient way (see Section 0). Particularly fast repetition is possible with minimal gradient switching as permitted by SPI (see Section 3.3) and ZTE (see Section 5.1). For these techniques, as an additional benefit, reducing TR
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helps approach the Ernst angle which may otherwise not be reached (see Section 5.5). Hence, in highbandwidth ZTE scanning, TR may range down to only hundreds of microseconds whereas UTE
7.2
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techniques are often operated with TR values of several or even tens of milliseconds. Gradient spoiling
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To obtain a strict steady state, the same net gradient area must be applied in all dimensions for each TR [120]. Hence, in radial sequences each readout gradient must be rewound and gradient spoilers
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must be applied along dimensions fixed in physical space to destroy unwanted coherences [38]. Therefore, despite sub-millisecond encoding times, TR often amounts to several milliseconds. To reduce TR, pseudo-steady-state operation can be employed by leaving out spoilers and instead randomizing gradient areas [65] and/or projection orders [7, 108]. The latter is common particularly in ZTE and SPI sequences where switching gradients is usually avoided to minimize TR (see Section 7.1) and to keep their near-silent nature. Spoiling is then performed by extending the readout gradient along with suitable projection ordering.
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7.3
Reducing scan time
3D radial scanning is relatively time-consuming, particularly at large matrix sizes. To fulfil the Nyquist criterion for center-out encoding, a factor of π more encoding steps are required compared with a conventional Cartesian sequence. Ways of reducing the scan time include generalized parallel imaging schemes [97, 121], angular undersampling [122] (see Section 7.4), and compressed sensing
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[123]. Furthermore, the k-space area covered per TR may be increased by using other center-out
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encoding schemes such as, e.g., spiral- [124], or cone-shaped [125-128] trajectories, within the limits given by gradient specifications and targeted T2. Another modification proposed for the UTE gradient
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shape aims at SNR optimization by adapting the k-space speed to compensate for the non-uniform
7.4
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density of radial encoding [129]. Geometry
The specific way of performing spatial encoding with the described techniques has implications for
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scan geometry.
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1) For both pure phase and radial frequency encoding there is no particular readout dimension in which bandwidth-limited filtering can be applied to restrict the FOV. Although possible for individual 1D projections in the radial techniques, such filtering would lead to inconsistent 3D data similar to selective RF excitation (see Section 5.5). Hence, in all (usually three)
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spatial dimensions a FOV given by object shape and the sensitivity range of the RF coils needs to be fully encoded to avoid aliasing. Thus potentially more encoding effort is required than with conventional sequences.
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2) As opposed to the discrete nature of the PSF in Cartesian imaging, the PSF for radial encoding exhibits a dense pattern of low-amplitude streaks. Therefore, undersampling causes streak artifacts of intensities much lower than with Cartesian sampling, and reducing the
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angular distance of the radial spokes by factors 2-4 leads to relatively benign and often
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tolerable artifacts.
3) For isotropic encoding, the FOV in 3D radial techniques is spherical. A simple adaptation to the object shape can be achieved by individual scaling of the gradients per dimension, resulting in an ellipsoidal FOV and anisotropic resolution. Disentangling these two parameters is possible with a generalized approach to radial encoding [130]. 4) Creating FOV off-centers with radial encoding is also more involved than with Cartesian techniques. For off-center excitation under a gradient, as in the SPI and ZTE sequences, the RF frequency must be adjusted to the respective gradient direction [7]. Similarly the receiver frequency needs to be set individually. In UTE encoding, due to gradient ramping, the receiver demodulation frequency must be changed also during the readout, which may be a
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limitation with some spectrometers. Alternatively, after sufficiently broadband acquisition, off-centers are applied by data demodulation during image reconstruction. 7.5
Contrast
The utility of medical images hinges on contrast between tissues types or conditions. MR image
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contrast is controlled by changing sequence parameters (e.g., excitation flip angle) or adding magnetization preparation (e.g., inversion), and further modified by combining multiple images (e.g.,
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by subtraction). In principle, in short-T2 imaging, the same concepts can be applied to generate
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contrast. However, a specific requirement is the separation of short-T2 signals from the often contrast generation in short-T2 techniques:
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dominant signals of long-T2 tissues. Furthermore, some specific considerations apply concerning
1) With negligible echo times, all tissues contribute according to their spin density and steadystate magnetization, often leading to images with relatively low intrinsic contrast.
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2) The effect of RF pulses on magnetization depends on their duration relative to T2 [52, 131] which needs to be taken into account for excitation, refocusing, and preparation. This may be
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limiting in some cases but can also be utilized to achieve T2 selectivity [76, 131-135]. 3) T2* weighting can be introduced in the UTE and SPI sequences by increasing TE, an option that obviously does not exist for the ZTE approach. A related effect may be achieved by
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reducing the encoding duration [7]. However, this approach emphasizes T2* decay at higher spatial frequencies.
4) T1 weighting using large flip angles is most feasible for the UTE technique but with long
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pulses, T2 decay during the pulse may need to be considered (see Section 4.3). For ZTE and SPI excitation, restrictions to flip angles exist associated with RF amplitude, dead time, and SAR (see Section 5.5).
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5) Most modes of magnetization preparation require more time than the TR values typical of the steady-state sequences described above. Hence, efficient performance requires a segmented
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implementation as employed in the Cartesian technique called “magnetization prepared rapid gradient echo imaging” (MP-RAGE) [136]. Emphasizing contrast by specific acquisition ordering as in the Cartesian case is not possible for basic UTE or ZTE scanning, as in all TR intervals equal parts of k-space are acquired. However, SPI and some of the hybrid techniques (see Section 0) enable such contrast manipulation. For the potentially SAR-intensive ZTE and SPI sequences, adding preparation pulses may be limited.
6) Subtraction images based on multiple gradient echo acquisitions with different TE can provide excellent T2* contrast. However, on radial trajectories, refocusing is sensitive to eddycurrent-induced artifacts due to concentration of errors in central k-space and requires
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correction measures [64, 66, 137-139]. For multi-echo ZTE scanning, eddy currents can be mitigated by employing the echo-shifting principle [140] to reduce gradient switching [141, 142]. 7) With zero or ultra-short TE, the image phase will be less affected by off-resonance than in conventional echo-based sequences. However, phase accrued during RF excitation and
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gradient encoding still contains information about local susceptibility or chemical shifts that
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can be used as a source of image phase contrast [143-145].
Within these constraints, contrast can be generated in short-T2 imaging. Several approaches have been
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proposed to separate short- from long-T2 signal, in particular to generate images with short-T2 signal only [38, 80, 132, 133, 146-158]. This results in so-called positive contrast for short-T2 tissues as
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opposed to classical T2-weighted images where these tissues are identified by the absence of signal (Fig. 18). Moreover, contrast within short-T2 components has been obtained by weighting by or quantification of T2 and T2* [53, 81, 159-163], T1 [162, 164-169], T1ρ [170, 171], chemical shift [150,
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172, 173], or magnetization transfer [167, 174-178]. Importantly, the last option requires suitable measures to separate true magnetization transfer from direct saturation which is stronger for lines
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broadened by rapid transverse relaxation [175]. 8 Artifacts and post-processing
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In this section, further aspects of image quality are discussed in addition to the resolution limitations and sequence-specific artifacts described in Sections 3-5. Off-resonance
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8.1
In radial images, the off-resonance effects due to chemical shifts, global field variations, or local susceptibility gradients deviate from the related signal shifts observed in Cartesian gradient echo
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techniques. With the center-out acquisition scheme often used for short-T2 MRI, off-resonance phase is accrued to form a cone-like shape in two-dimensional k-space or a corresponding structure for the
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3D case. This leads to blurring or ringing depending on the resonance offset. The transition between the two effects occurs for offsets between 1 and 2 PWB, where the pixel bandwidth of an image is defined as PBW = BW / N and N is the number of pixels in the FOV. To reduce off-resonance due to B0 non-uniformity, shimming is usually applied. The generic means of
mitigating off-resonance artifacts of any origin is using strong gradients, i.e., large PBW, as employed anyhow for encoding short-T2 tissues. Complementary, off-resonance correction can be applied to the data based on B0 mapping, using conjugate phase reconstruction [179-182] or iterative approaches [183, 184]. As in other sequences, a range of off-resonances values within a voxel leads to signal loss
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until the start of data acquisition and signal decay during the readout, which can only be mitigated by shortening TE and increasing encoding gradients, respectively. A particular effect in vivo is created by the chemical shift between water and fat of approximately 3.5 ppm. This corresponds to hundreds of Hertz at usual field strengths which is on the order of commonly used PBWs. If both water and fat contribute to the MR signal, the demodulation frequency
F
chosen for data acquisition and/or image reconstruction can be on resonance for only one of them,
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leading to image blurring for the respective other component. This has a strong impact on the appearance of tissue interfaces and thus image contrast as demonstrated in Fig. 19. It is important to
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keep these effects in mind in practice because artifacts such as those seen in the figure might be misinterpreted. Besides using higher PBW, such effects may be avoided by suppressing either fat or
8.2
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water using magnetization preparation or separation techniques [144, 185-190]. Resolution enhancement
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As outlined in Section 3.1, resolution in MRI is limited by apodization in k-space caused by T2 decay affecting the acquisition range. However, in analogy to time-domain corrections in spectroscopy
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[191], such apodization can be countered by suitable correction factors in k-space, corresponding to a filter operation in the image domain. In particular, T2 filtering can be applied which compensates for exponential T2 decay (Fig. 20):
',!- &
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,!- &
(5)
When setting ',!- & of a given target tissue, the spatial resolution of this tissue will not be
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limited by the encoding duration. However, this tempting approach has two serious limitations: 1) While resolution will be improved for all . ',!- &, high spatial frequencies will be overemphasized for any / ',!- &, leading to ringing artifacts. For samples with excessive
U
T2 range, the spatial variation of T2 can be included in the signal model [192]. However, this
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approach requires additional T2 mapping and more extensive computation.
2) Multiplication by the exponential in Eq. (5) boosts not only high-k signal but also noise, at the expense of image SNR.
Therefore, although moderate T2 filtering may be a useful means of optimizing image appearance, it cannot replace sufficiently rapid encoding in short-T2 MRI. 9 Hardware As discussed in preceding sections, imaging of tissues with short T2 differs from common MRI in terms of acquisition strategies and sequence timing. As a consequence, it alters requirements on the
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instrumentation used. In several respects, short-T2 imaging calls for enhanced hardware performance while in others it demands lower or different specifications. For some short-T2 techniques, dedicated hardware is not essential. Nevertheless, for best results and for conquering ever shorter T2, tailored hardware is clearly beneficial if not indispensable. Magnet
F
9.1
Today, the strength of the static B0 field of most clinical MRI systems is 1.5 T or 3 T, whereas in
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numerous research systems it is 7 T and reaches up to 10.5 T [193]. B0 affects mainly three aspects:
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1) Baseline sensitivity: As usual in human MRI setups, the SNR increases approximately linearly with B0 [194]. SNR is particularly critical for tissue components with short T2 due to
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often low spin density and long T1 (see Section 2.5) as well as short, high-bandwidth acquisitions. Thus, with regard to baseline sensitivity, high B0 is favored. 2) Relaxation times: For field strengths and tissues of interest, T1 increases along with B0 while
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both T2 and T2* decrease [13], eating up part of the sensitivity gain. In turn, at high field, shorter T2 calls for potent short-T2 techniques and longer T1 entails lower optimal flip angles.
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3) B1 efficiency: The power efficiency of generating B1 declines as B0 increases, approximately according to 0 120 0 [195, 196] with 120 being the power absorbed in the exposed tissue. This dependency renders RF power demands and SAR at high field challenging
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despite slightly reduced flip angle requirements.
Optimal choice of the field strength accounts for all of these aspects and also depends on the tissue
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and application of interest.
Concerning B0 uniformity, short-T2 techniques pose reduced demands since they use strong gradients and high-bandwidth acquisition. Therefore, per-subject shimming is often not required. It still tends to
U
be relevant, however, when weaker or ramped gradients are employed during data acquisition, such as in WASPI, UTE imaging, or gradient-modulated variants (see Section 6.2). Furthermore,
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magnetization preparation may require more uniform B0 than data acquisition (see Section 7.5). 9.2
Gradient system
The gradient system [197] has two main types of components, amplifiers and coils, which together produce dynamic field gradients according to input waveforms. 1) Gradient strength: The maximum gradient strength G is a key parameter as it determines kspace speed, which competes with signal decay. For purely frequency-encoded techniques (e.g. UTE or ZTE imaging) the latter causes approximately exponential apodization (Figs. 8 and 17). To actually achieve a given nominal resolution %&, the readout gradient must
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approximately obey 3 $ %& [37, 38]. For example, for T2 = 500 μs an actual resolution of 1 mm can be obtained with G = 23.5 mT/m, which is feasible on most clinical systems, providing G is in the range 20 - 80 mT/m. However, T2 = 50 μs requires the extreme value of G = 235 mT/m or will yield only 10 mm resolution at the lower G. For purely phaseencoded techniques (e.g. SPI), resolution is not limited by T2, but large G is still important to
F
limit TE for best SNR (see Section 3.3 and Fig. 17).
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2) Slew rate: Techniques that set readout gradients before RF excitation (e.g., SPI and ZTE) may operate gradients quasi-continuously with minimized switching (see Section 5) and thus do
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not depend on high slew rates. Still, faster gradient switching permits shorter TR, which is associated with smaller optimal flip angle, higher acquisition duty cycle, and less overall scan
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time (see Section 7.1). In contrast, when RF excitation is performed first (e.g. in UTE), the speed of gradient ramping is a key determinant of resolution and robustness against offresonance (see Section 4.1, Fig. 8). Slew rates in commercial scanners range up to 200
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mT/m/ms in combination with a gradient strength of, e.g., 40 mT/m. Importantly, to efficiently use an increase in gradient strength, the slew rate must grow quadratically [119].
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3) Duty cycle: Quasi-continuous operation of ZTE- and SPI-like sequences means that gradients are always on, often at constant net magnitude, but with changing direction. This can only be accomplished if full duty cycle is provided at the applied high strength, which is a rather
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uncommon design criterion for gradient systems. Although the same techniques can also be run non-continuously with gradient switching and reduced duty cycle, the continuous mode is clearly preferable as it permits shorter TR and strongly reduces eddy currents and acoustic
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noise. For methods with intrinsic gradient switching (e.g. UTE imaging), high duty cycle is less essential but still helps minimize TR.
4) Amplifier: Requirements for the gradient amplifier arise from the three specifications
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discussed above. Gradient strength and slew rate are directly related to maximum current and voltage, respectively. Some amplifiers offer added flexibility by permitting multiple
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combinations of peak current and peak voltage [198]. Operation at full duty cycle requires that the current necessary for a desired gradient strength can be provided continuously along with the corresponding power for given ohmic losses in the gradient coil.
5) Fidelity: The spatiotemporal fidelity of gradient fields is generally hampered by delays, finite bandwidth, eddy currents, and amplifier imperfections. This affects image quality primarily in UTE techniques (see Section 4.2) and when additional echoes are generated for creating contrast (see Section 7.5). To a certain degree, these effects can be reduced by delay calibration and pre-emphasis. If necessary, residual field errors can be measured and accounted for at the level of image reconstruction (see Section 9.5).
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6) Peripheral nerve stimulation: In vivo, gradient switching must avoid peripheral nerve stimulation (PNS) induced by rapid, large-amplitude field changes [199]. In this respect, short, localized gradient coils are advantageous over whole-body gradients [197]. Short-T2 techniques with minimized gradient switching are not a problem with regard to PNS. Sequences with switching need more attention but are typically less prone to PNS than, e.g.,
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echo-planar imaging with long trains of shortly spaced rapid field changes.
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In summary, fulfilling all of the listed criteria at the same time is not trivial but it is possible with
9.3
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dedicated designs, thus also enabling extreme applications [200]. RF chains
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Short-T2 MRI relies on high-bandwidth, sub-millisecond acquisitions and thus also alters requirements on RF hardware. For UTE imaging, the factors to consider are rapid T/R switching and avoiding background signal. For ZTE and SPI techniques, which require excitation during readout
power is needed (see Section 5.5).
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gradients, these demands are further increased, and in addition, efficient RF transmission at high
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1) Amplifier: For ZTE and SPI, the bandwidth of RF amplifiers must be sufficiently large to enable high-bandwidth excitation using either short hard pulses of duration in the μs range or frequency-swept pulses with large sweep width (see Section 5.5). Moreover, rapid amplitude
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changes at the beginning and end of these pulses call for sub-microsecond rise times. In analogy to gradient amplifiers, attenuation in transition bands may be overcome by preemphasis as far as permitted by power margins. Moreover, accurate realization of modulated
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pulses requires good linearity which may also be improved by means of calibrated compensation. High-power transmission is particularly important to achieve adequate flip angles by hard-pulse excitation. As an example, when using an RF coil providing a typical B1
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efficiency of 20 45 67, obtaining a flip angle of 5° with a hard pulse of 2 μs duration
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requires 2.9 kW. Furthermore, the available duty cycle must permit application of the desired pulses (few to hundreds of microseconds) at short TR (potentially sub-millisecond, see Section 7.1) and for long scan times (potentially tens of minutes). Amplifier ring-down after the RF pulse causes artifacts in high-bandwidth images when time constants are similar to T/R switching times [201]. To suppress this ring-down sufficiently, blanking after transmission must be fast and achieve a high degree of isolation in addition to the attenuation provided by T/R switching or de-tuning (see below).
2) Coils: MRI generally relies on either a single coil for both RF transmission and detection or separate devices for the two purposes. The latter setup usually consists of a receiver array
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near the subject to maximize SNR and a surrounding transmit coil. However, those short-T2 techniques with high B1 requirements call for particularly high transmit efficiency, which can favor T/R coils, in particular T/R arrays. A further advantage of such single-coil setups are reduced space requirements along the radial direction. On the other hand, for safety reasons, the distance to the subject needs to be larger than for a receive-only coil. Furthermore, pre-
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amplifier decoupling cannot be implemented. On the receive side, as usual, coil arrays offer
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SNR benefit [202] and parallel imaging capability [97]. In ZTE techniques, complementary spatial encoding by array detection also holds potential to support filling the k-space gap (see
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Section 6.1). Similar to RF amplifiers, the bandwidth of RF coils must be large enough to generate and detect, at sufficient uniformity, suitable excitation pulses and resonance signals,
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respectively. If not provided by conventional coils, dedicated designs are required. For residual imperfections, appropriate pulse compensation or data correction procedures may be employed. A further design criterion arises from the original aim to capture short-T2 signals
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from the subject. In addition to the latter, short-T2 detection is prone to capture MR signals also from hardware parts, which conventional MRI does not pick up due to their rapid decay.
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This issue is most prominent for materials close to coil conductors where B1 is high (see Section 5.4). Ideally, such contributions are avoided by choosing materials that do not contain the targeted nucleus. For 1H imaging this is achieved with, e.g., glass or fluorinated polymers such as polytetrafluoroethylene (PTFE) [82, 203-205].
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3) T/R switching: In pulsed MR, the receive chain must be sufficiently isolated from the transmit path in both operation modes. During transmission, destruction and preferably even saturation
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of receive pre-amplifiers must be prevented. During reception, residual ring-down of the power amplifier must be blocked. Depending on the coil configuration, either switching between the two operation modes by means of a T/R switch or de-tuning of the transmit coil and tuning of the receive coil are required. The particular challenge for short-T2 techniques is
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to perform this rapidly, i.e., in the low microsecond range as opposed to several tens of
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microseconds in commercial systems, to minimize the RF dead time [206, 207] (see Section 5.2). To avoid image artifacts, switching transients must have decayed to a level well below that of the MR signal when data is acquired. Rapid switching is particularly challenging at high peak and average power. It is typically based on positive-intrinsic-negative (PIN) diodes [208], but has also been achieved with transistors [208, 209] or micro-electro-mechanical systems (MEMS) [210]. In a recent design based on PIN diodes, switching transients are cancelled by symmetric biasing [211].
4) Simultaneous operation: A tempting alternative to RF switching is simultaneous transmission and reception [212-214]. With this approach, dead time overheads are avoided and sequences
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are simplified. Simultaneous transmission and reception was common practice before the advent of pulsed NMR, albeit in rigid setups and at much lower frequencies. For in-vivo imaging at present-day MR frequencies, sufficient isolation between transmit and receive paths is difficult to achieve without T/R switching or detuning. Isolation between the two RF paths can be achieved by geometric coil decoupling, additional decoupling by passive
F
electronics, and active cancelation of transmit leakage into the receive path. However, the
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robustness of such setups is limited by residual variation in RF conditions, particularly by patient motion, which causes subtle changes in RF loading. Feedback control of cancelation
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has been proposed to address this limitation [215, 216]. In an alternative approach, isolation is facilitated by sideband modulation, which permits performing transmission and reception in
9.4
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different frequency bands [217]. Console
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The console of an MRI system includes a sequencer generating RF and gradient waveforms, a receiver for data acquisition, and a reconstructor that calculates images from raw data. This
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architecture is the same for short-T2 imaging. However, for high performance in the short-T2 mode, the components should meet more stringent specifications. 1) Sequencer: The temporal resolution of the sequence description should be 1 μs or below. The
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definition and generation of RF pulses must be more flexible than in common clinical MRI systems, including a wider range of pulse durations (1 μs to tens of milliseconds) and TR (down to 100 μs, see Section 7.1). To perform off-center imaging with ZTE or SPI (see
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Section 7.4), per-shot adjustment of the excitation carrier frequency is necessary. The same is true for the receiver with any radial technique.
2) Receiver: The feasible acquisition bandwidth must cover the signal bandwidth of up to the
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megahertz range. Even higher acquisition bandwidth is needed for oversampling to enable finite-support extrapolation (see Section 5.3). In conjunction with high spatial resolution and
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array detection, 3D acquisition with oversampling yields very large amounts of data that need to be handled and stored. High bandwidth and short TR render these data challenging also in terms of throughput. Short-T2 imaging benefits particularly from flexibility at the level of digital filtering, where short, customized, and multi-rate filters [218] permit minimizing RF dead times without bloating the raw data (see Section 5.2).
3) Reconstructor: 3D radial data sets tend to be very large and do not readily lend themselves to splitting into smaller portions for separate processing. Therefore, a computer used for image reconstruction should be equipped with ample memory to avoid time-consuming swapping on the hard disk. For example, the raw data for image matrix size 256 acquired with full Nyquist
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angular sampling, 4-fold radial oversampling, and 8 channels amount to approximately 13 GB. For 3D gridding reconstruction, about 8 times that amount of memory needs to be allocated. Reconstruction time is longer than in standard Cartesian imaging and is determined by the numerical complexity of algebraic parts, density correction, gridding interpolation, and FT, and by the number of cycles in the case of iterative reconstruction (see Sections 4.5 and
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5.3). For speed-up, reconstruction may be partly parallelized and distributed to different nodes
Calibration tools
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9.5
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on a single computer or a cluster of computers, or on graphical processing units (GPUs).
All magnetic fields – static, audio- and radio-frequency – experienced by the spins contribute to
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encoding of the MR signal. Hence, for proper image reconstruction, the spatial and temporal behavior of these fields must be accurately included in the underlying signal model. However, encoding fields are produced with finite accuracy by the respective hardware, i.e., the magnet, gradient subsystems,
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and RF chains. To determine actual fields and field evolutions, either the MR system itself or additional dedicated measurement hardware may be used to record field information at runtime or in a
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calibration step. Depending on the severity of deviations, the results can be used to improve image reconstruction or to pre-distort input waveforms for higher accuracy. 1) Gradients: Apart from the methods discussed in Section 4.2, spatial and temporal behavior of
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gradient fields per se as well as eddy current effects can be measured by a set of MR field probes [71]. A field probe is a small RF coil wrapped around an NMR sample and determines the variation over time of field strength at one position. Several such probes enable spatial
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expansion of the field dynamics in the imaging volume. Such information has been used to improve reconstruction of, e.g., UTE images [219, 220] and dual-echo ZTE data [139]. Instead of measuring field behavior for each protocol, by assuming linearity and time
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invariance, the gradient system may also be characterized by its impulse response [221], enabling field calculations for arbitrary gradient shapes.
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2) RF: Effective RF waveforms are subject to variable degrees of distortion caused by RF amplifiers, lines, cables, and coils. These are most easily observed with a pick-up coil placed near the sample and connected to an MR spectrometer or a spectrum analyzer. In this way, effective RF pulses as experienced by the spins are observed, including all disturbing effects in the transmit chain. Alternatively, field probes may be used in an advanced setup to monitor RF and gradient fields simultaneously [222], including information on spatial variation. Knowledge of effective RF waveforms is particularly important for the use of amplitude- or phase-modulated pulses (see Section 5.5).
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9.6
Summary
The hardware considerations compiled here give an indication of how dedicated systems for short-T2 MRI might differ from present-day instrumentation. For optimal UTE imaging, requirements are actually rather similar to those of conventional MRI. High B0 should be combined with high gradient performance in terms of amplitude, slew rate, and
F
fidelity, and careful consideration of PNS. With regard to RF, in a similar manner to regular MRI,
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dense receiver arrays are desirable for optimal SNR. As one feature specific to the short-T2 scenario, a UTE system should offer rapid T/R switching to minimize TE. Integrated field monitoring equipment
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would facilitate accurate gradient characterization.
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In contrast, a system designed for ZTE imaging would differ more fundamentally. Due to the more limiting role of RF power considerations, the optimal field strength will be lower for ZTE imaging than for UTE. B0 field homogeneity is less critical in the ZTE scenario and may be traded off against
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magnet length and cost. Gradients should offer high amplitude and, critically, unlimited duty cycle. However, neither high slew rate nor gradient fidelity are particularly important, which will alter and
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potentially simplify coil design and amplifier considerations. Likewise, PNS is of little concern. On the RF side, the ZTE mode is more demanding, calling not only for dense array detection but also for high-power transmit capability, efficient transmission, very rapid T/R switching, and 1H-free coil front-ends. A ZTE console should offer superior temporal definition, whereas integrated RF
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monitoring would improve waveform correction and thus expand the range of feasible pulse modulation schemes.
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10 Applications
Most applications of short-T2 MRI relate to its ability to depict solid tissues. This is in contrast to conventional MRI, which depicts such tissues mostly by the absence of signal, relying on so-called
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negative contrast. By using dedicated short-T2 techniques instead, signals from solid tissues can be
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observed directly and thus with positive contrast. This approach has two major advantages: 1) Short-T2 tissues can be distinguished from actually void spaces.
2) Signal variation related to spin density or other sources of contrast (see Section 7.5) permits differentiation between different short-T2 tissues or between healthy and pathological tissues.
Table 2 lists several tissues and materials of interest along with T2 values as reported in the literature, indicating fields of application. With conventional MRI hardware, the short-T2 techniques described previously enable imaging of T2 values with several hundreds of microseconds at useful resolution. By employing dedicated RF and gradient hardware (see Section 9), T2 values below 100 μs also become realistic targets [201, 223].
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1) Musculoskeletal (MSK) system: Short-T2 tissues are particularly prevalent in the MSK system, such as bone, tendons, ligaments, menisci and cartilage (Fig. 21a). The focus of many investigations is to image bone and quantify bone mineral density, either indirectly from bone water [49, 51, 53, 111, 151, 160, 168, 172, 176, 224-246] or directly from hydroxyapatite employing
31P-imaging
[6, 22, 244, 247-257], due to its potentially high diagnostic value,
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e.g., in osteoporosis. This application has also been expanded to bone substitutes [258].
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Furthermore, tendons [31, 38, 150, 170, 259-261], ligaments [49, 54], menisci [49, 51, 150, 170, 262], and cartilage [263-265] are of specific interest due to frequent injuries,
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inflammations, or wear.
2) Positron emission tomography (PET): The ability to depict bone can be utilized in devices for
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combined PET and MRI. PET requires attenuation correction for bone and other tissues which may be based on segmentation of short-T2 MR images [266-273]. The same approach can be applied for attenuation caused by hardware parts of the MRI scanner, in particular RF
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coils.
3) Pulmonary MRI with short-T2 techniques [47, 99, 108, 162, 167, 274-301] (Fig. 21b) has
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large potential, e.g., to improve specificity in early diagnosis of certain diseases (e.g., fibrosis) as compared to computer tomography (CT). For this application, a particular advantage of the radial short-T2 techniques is high robustness with respect to motion. 4) Iron oxide nanoparticles (IONP) are used as contrast agents for, e.g., cell labeling or tracking
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drug delivery. Typically, in MR images IONPs are observed as signal voids due to reduced T2*. Alternatively, short-T2 techniques enable imaging them with positive contrast, thus
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utilizing reduced T1 instead of T2* [7, 149, 165, 302-306]. In this way, their identification and localization can be improved.
5) Dental MRI using short-T2 techniques allows depiction of teeth with high quality (Fig. 21c). Application in vivo holds promise for improved early diagnosis of, e.g., caries lesions and
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benefits from dedicated hardware [307-315]. However, at present patient comfort, scan times,
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and cost cannot yet compete with standard X-ray techniques.
6) Myelin is a biomaterial that is essential in rapid electrical signal propagation along nerve fibers in white matter of the brain and the spinal cord. Therefore demyelination plays an important role in numerous neurological disorders. In a conventional approach, myelin is quantified indirectly via myelin water [316, 317]. Short-T2 techniques have the potential to directly visualize and quantify myelin but must rely on effective suppression of long-T2 tissue components [35, 147, 318-325].
7) Plaques in vessels due to calcification by arteriosclerosis reduce or obstruct blood flow in, e.g., peripheral, cardiac, or brain vasculature. Due to the short T2* of plaques, their assessment
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is conventionally based on X-ray imaging. Short-T2 MRI offers a prospective alternative modality for plaque detection and specification [319, 326-329]. 8) Functional MRI (fMRI): MR techniques with ultra-short or zero TE are not intrinsically sensitive to blood-oxygenation-level dependent (BOLD) contrast. Nevertheless, they can be used for performing fMRI. Sensitivity to neuronal activation is either obtained by T2
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preparation [330] or based on other mechanisms, e.g., perfusion [331, 332]. As a particular
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advantage, the zero-TE techniques offer silent operation, avoiding brain activation related to acoustic noise generated by gradient switching.
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9) Electric property tomography (EPT): Spatial mapping of electric permittivity and conductivity by EPT has potential for various applications [333]. EPT is based on evaluation
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of the image phase. Techniques with negligible TE may contribute in this field due a reduced sensitivity of their image phase to off-resonance and eddy currents [334]. 10) Sodium MRI: Sodium is of interest due to its important role in many biochemical processes
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and is a potential biomarker in the diagnosis of diseases such as stroke, cancer, and osteoarthritis. The NMR-active isotope 23Na has spin 3/2, leading to bi-exponential relaxation
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behavior (see Section 2.3). T2 of the shorter component is typically best addressed with dedicated short-T2 techniques [126, 129, 335, 336]. 11) Materials: Apart from in-vivo uses, short-T2 MRI has large potential for materials research [337-340] and inspection as well as for applications in the food industry [341-343]. The
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ability to depict short-T2 materials is also instrumental in the non-destructive examination of mummies [344].
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12) Inhomogeneous B0: The use of minimized TE and high bandwidth, short-T2 techniques also provides considerable resilience against signal loss and displacement from strong macroscopic B0 (see Section 2.6). This kind of robustness improves imaging in the presence
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of tissue-air interfaces or metal implants [7, 345].
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11 Conclusions
In this article, we have reviewed the growing family of techniques for in-vivo MRI of tissues with rapid decay of transverse magnetization. Besides covering the variety of current approaches and implementations, this survey has taken a conceptual perspective, emphasizing fundamental challenges, derived principles, as well as commonalities and key differences among the various techniques. It has been shown that, despite their variety, short-T2 imaging techniques can be categorized according to a few fundamental design choices, leading to four basic sequence ideas, CTI, SPI, UTE, and ZTE. The Cartesian techniques CTI and SPI offer the attractive feature of enabling theoretically unlimited
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spatial resolution even in the case of rapid signal decay. However, this option comes at the expense of poor SNR. Greater SNR can be achieved with the radial UTE and ZTE sequences, which are most widely used at this point. UTE imaging is feasible with conventional hardware but requires accurate gradient calibration. In comparison, ZTE approaches offer higher short-T2 sensitivity and resolution but are more demanding in terms of RF hardware and image reconstruction. The choice of technique
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thus depends on the targeted application and technical resources. Offering guidance for these
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considerations is one of the chief purposes of this article.
As an outlook, two roads can be foreseen in the emerging field of short-T2 MRI. First, clinical
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usefulness needs to be investigated for applications that are readily accessible with existing commercial scanners. These relate to tissues with T2 above approximately 300 μs, such as in MSK
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and pulmonary imaging. The basis for this task will be relatively established implementations of both UTE and ZTE variants, whereas further effort will be required to verify reliable contrast options. Second, the feasibility of imaging in vivo at T2 values below a few hundreds of microseconds,
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possibly down to tens of microseconds (as, e.g., in myelin) still needs to be explored and requires additional methodological and technical efforts. With regard to imaging methods, it is expected that
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there is still considerable potential in fusing and optimizing current approaches. Furthermore, on the technical side, advances in high-bandwidth capabilities of MR scanner hardware with respect to gradient strength and duty cycles, RF switching speed, console specifications, and proton-free
12 Appendices
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materials will pave the way for further discoveries in the world of short-T2 tissues.
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12.1 Implementation of the rho filter
For sequences with radial encoding, a suitable DCF is efficiently obtained by calculation of the rho filter. For center-out encoding schemes as used in UTE and ZTE imaging, it may be implemented as
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follows:
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89
:
89! ! !
(A1) :
! !
89! ! !
:
:
! !
(A2) :
(A3)
Here, ! ;;! counts the radial k-space positions starting from the center, k-values are given in arbitrary units, and d is the dimensionality of the data. Note that the DCF calculated according to Eqs. (A1-A3) is valid also for not being at as well as non-equidistant k-space positions.
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12.2 Steady-state magnetization at short TR In a steady-state sequence with complete spoiling of transverse magnetization after each TR interval, the flip angle for maximized transverse steady-state magnetization ' is the Ernst angle [118] ( <=>>?@
(A4)
F
with )
(A5)
O
O
leading to [346]
'
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For ) * the approximation [119] )
E-
#
( # &AABC
)
#
as well as
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leads to
)
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' #
)
(A6)
(A7)
(A8)
(A9)
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which by neglecting the -1 further simplifies to ' #
)
(A10)
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Hence at very short TR, both the Ernst angle and the signal are approximately proportional to ).
Funding
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This research did not receive any specific grant from funding agencies in the public, commercial, or not-for-profit sectors. Declarations of interest MW None
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Declarations of interest KP
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Klaas Pruessmann holds a research agreement with and receives research support from Philips Healthcare. He is a shareholder of Gyrotools LLC and Skope
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Magnetic Resonance Technologies.
References
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[1] P. Mansfield, Pulsed NMR in solids, Prog Nucl Mag Res Sp, 8 (1971) 43-101. [2] E.R. Andrew, E. Szczesniak, A historical account of NMR in the solid state, Prog Nucl Mag Res Sp, 28 (1995) 11-36.
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[3] P. Jezzard, J.J. Attard, T.A. Carpenter, L.D. Hall, Nuclear magnetic resonance imaging in the solid state, Prog Nucl Mag Res Sp, 23 (1991) 1-41. [4] J.B. Miller, NMR imaging of materials, Prog Nucl Mag Res Sp, 33 (1998) 273-308. [5] S. Emid, J. Creyghton, High resolution NMR imaging in solids, Physica B+ C, 128 (1985) 81-83.
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[366] K.S. Nayak, J.M. Pauly, G.E. Gold, D.G. Nishimura, Imaging ultra-short T2 species in the brain, In Proceedings of the 8th Annual Meeting of ISMRM, Denver, USA, 2000, p.509
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[367] E. Ercan, P. Boernert, A. Webb, I. Ronen, Whole-brain tissue-based assessment of the ultrashort T2 component using 3D UTE MRI relaxometry, In Proceedings of the 20th Annual Meeting of ISMRM, Melbourne, Australia, 2013, p.4279 [368] Q. He, Y. Ma, S. Fan, H. Shao, V. Sheth, G.M. Bydder, J. Du, Direct magnitude and phase imaging of myelin using ultrashort echo time (UTE) pulse sequences: A feasibility study, Magn Reson Imag, 39 (2017) 194-199.
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[369] S.J. Fan, Y. Ma, E.Y. Chang, G.M. Bydder, J. Du, Inversion recovery ultrashort echo time imaging of ultrashort T2 tissue components in ovine brain at 3 T: a sequential D2O exchange study, NMR in biomedicine, 30 (2017) e3767. [370] S.J. Fan, Y. Ma, Y. Zhu, A. Searleman, N.M. Szeverenyi, G.M. Bydder, J. Du, Yet more evidence that myelin protons can be directly imaged with UTE sequences on a clinical 3T scanner: Bicomponent T2* analysis of native and deuterated ovine brain specimens, Magn Reson Med, 80 (2018) 538-547. [371] A. Ramani, A. Aliev, G. Barker, P. Tofts, Another approach to protons with constricted mobility in white matter: pilot studies using wideline and high-resolution NMR spectroscopy, Magn Reson Imag, 21 (2003) 1039-1043. [372] M.A. Schmidt, G.Z. Yang, P.D. Gatehouse, D.N. Firmin, FID-based lung MRI at 0.5 T: theoretical considerations and practical implications, Magn Reson Med, 39 (1998) 666-672. [373] M. Kveder, I. Zupančič, G. Lahajnar, R. Blinc, D. Šuput, D. Ailion, K. Ganesan, C. Goodrich, Water proton NMR relaxation mechanisms in lung tissue, Magn Reson Med, 7 (1988) 432-441.
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[374] S.M. Triphan, F.A. Breuer, D. Gensler, H.U. Kauczor, P.M. Jakob, Oxygen enhanced lung MRI by simultaneous measurement of T1 and T2* during free breathing using ultrashort TE, J Magn Reson Imag, 41 (2015) 1708-1714. [375] E.D. Pracht, J.F.T. Arnold, T.T. Wang, P.M. Jakob, Oxygen-enhanced proton imaging of the human lung using T2*, Magn Reson Med, 53 (2005) 1193-1196.
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[376] H. Hatabu, D.C. Alsop, J. Listerud, M. Bonnet, W.B. Gefter, T2* and proton density measurement of normal human lung parenchyma using submillisecond echo time gradient echo magnetic resonance imaging, Eur J Radiol, 29 (1999) 245-252.
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[377] R.J. Theilmann, T.J. Arai, A. Samiee, D.J. Dubowitz, S.R. Hopkins, R.B. Buxton, G.K. Prisk, Quantitative MRI measurement of lung density must account for the change in T2* with lung inflation, J Magn Reson Imag, 30 (2009) 527-534.
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[378] N. Beckmann, B. Tigani, L. Mazzoni, J.R. Fozard, MRI of lung parenchyma in rats and mice using a gradient-echo sequence, NMR in biomedicine, 14 (2001) 297-306.
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[379] A. Zhu, D. Hernando, K.M. Johnson, S.B. Reeder, Quantification of short-T2* signal components in the liver using radial 3D UTE chemical shift-encoded MRI, In Proceedings of the 25th Annual Meeting of ISMRM, Honolulu, Hawaii, USA, 2017, p.120
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[380] E.K. Doyle, K. Toy, B. Valdez, J.M. Chia, T. Coates, J.C. Wood, Ultra-short echo time images quantify high liver iron, Magn Reson Med, 79 (2018) 1579-1585.
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[381] S. Codd, M. Mallett, M. Halse, J.H. Strange, W. Vennart, T. Van Doorn, A three-dimensional NMR imaging scheme utilizing doubly resonant gradient coils, J Magn Reson B, 113 (1996) 214-221.
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[382] M. Fernandez-Seara, S. Wehrli, F. Wehrli, Multipoint mapping for imaging of semi-solid materials, J Magn Reson, 160 (2003) 144-150.
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Glossary 1D: one-dimensional 3D: three-dimensional AC: autocorrelation
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AR: algebraic reconstruction
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B0: static magnetic field
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B1: radio-frequency magnetic field
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BLAST: back-projection low angle shot BOLD: blood-oxygenation-level dependent BW: bandwidth
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CT: computer tomography
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CTI: constant time imaging DCF: density correction function
EPT: electrical property tomography
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FID: free induction decay
fMRI: functional magnetic resonance imaging
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FOV: field-of-view
FT: Fourier transform
GM: gradient-modulated
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GPU: graphical processing unit
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HYFI: hybrid filling IONP: iron oxide nanoparticle MEMS: micro-electro-mechanical system MP mapping: multipoint mapping MP RAGE: magnetization prepared rapid gradient echo imaging MR: magnetic resonance MRI: magnetic resonance imaging
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
MSK: musculoskeletal MTF: modulation transfer function NMR: nuclear magnetic resonance PBW: pixel bandwidth
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PET: positron emission tomography
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PETRA: pointwise encoding time reduction with radial acquisition
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PIN: positive-instrinsic-negative
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PNS: peripheral nerve stimulation PSF: point spread function PTFE: polytetrafluoroethylene
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RF: radio-frequency
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RHE: ramped hybrid encoding RUFIS: rotating ultra-fast imaging sequence SAR: specific absorption rate
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SPI: single-point imaging
SPRITE: single-point ramped imaging with T1 enhancement
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SNR: signal-to-noise ratio
SWIFT: sweep imaging with Fourier transformation T1: longitudinal relaxation time
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T2: transverse relaxation time
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T2*: apparent transverse relaxation time TE: echo time : acquisition range, or range of times across which data are collected
TR: repetition time T/R: transmit-receive UTE: ultra-short echo time WASPI: water- and fat-suppressed proton projection imaging
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
ZTE: zero echo time
Fig. 1: Dipolar coupling as an example of an orientation-dependent interaction. a) The dipolar field of one spin (red arrow) affects the total field at the location of another spin (blue arrow). The local field contribution %0E >?@ E depends on the orientation of the molecule as the spins keep their
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alignment with B0 upon molecular rotation, thus changing the position of the affected spin within the
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dipolar field. The relevant angle θ is the angle between the B0 field vector and the axis connecting the two spins (dashed line). Molecular tumbling changes the orientation of the molecule over time and the
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local field as seen by the spin fluctuates. b) Effects of dipolar coupling on NMR spectra. In a single crystal, all local lattices form the same angle with B0 (θ = 20° in this case), leading to a distinct peak.
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The second, symmetric peak arises from spin pairs in the same sample with anti-parallel instead of parallel mutual alignment (one of the spins in a) flipped). In a powder, all lattice orientations co-exist; thus the spectrum reflects all corresponding frequency shifts. The latter are weighted with @FG E
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according to their occurrence for an equal distribution of orientations, leading to the peculiar shape peaking at θ = 90°, the so-called Pake pattern. Again, both ways of alignment contribute. Note that
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the field shift vanishes at the so-called magic angle of θ ≈ 55°. In a hypothetical liquid with arbitrarily fast molecular rotation, the dipolar field is averaged to zero over time, leading to a single narrow line. In a real liquid, averaging also occurs, but in addition, the temporal fluctuations cause T1 and T2 relaxation, the latter leading to line broadening. Note that the spectra are scaled for presentation
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purposes. c) Free induction decay signals obtained for the spectra shown in b). The hypothetical liquid gives a constant on-resonant signal whereas the liquid with relaxation shows exponential decay. The
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off-resonant peaks of the single crystal lead to oscillations with constant envelope. The signal associated with the powder spectrum drops rapidly due to superposition of a wide range of frequencies. In this case, the observed loss of transverse magnetization is a non-random, reversible
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process and therefore not considered to be actual relaxation. Fig. 2: Relaxation in a nutshell. In the presence of an orientation- or position-dependent interaction
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(e.g. dipolar coupling, see Fig. 1), molecular tumbling leads to randomly fluctuating fields ΔB(t) at the spins’ locations. The temporal behavior of the random field is determined by the speed of the tumbling motion and characterized by the correlation time τc, whereas its amplitude depends on the strength of the interaction. Two typical random fields are shown with rapid (τc = 0.05 ns, red) and
slow fluctuations (τc = 0.4 ns, blue). The autocorrelation (AC) function calculated from the average of such a random field decays exponentially with the constant τc. Consequently, the Fourier transform (FT) of the AC function is a Lorentz function, also known as the spectral density J(ω) of the random field, i.e., the energy available at different frequencies, which will be required below. Relaxation of longitudinal magnetization involves changes of the populations of different spin states. This is
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
illustrated in the energy diagram drawn for an isolated pair of equivalent spins, where transitions between combined states occur with probabilities W0, W1, and W2 where the subscripts indicate the corresponding “step size” in the diagram. The transition probabilities scale with the spectral density taken at multiples of the Larmor frequency ω0. How they contribute to T1 is given by the particular interaction causing the relaxation. T1 relaxation is most efficient for τc ≈ 1/ω0. Transverse relaxation
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involves losses in coherences. Therefore, J(0) contributes and dominates T2 for large τc, leading to a
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continuous T2 reduction for slower motion. The dashed vertical red and blue lines indicate the relaxation times obtained for the example values of τc.
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Fig. 3: Principles of short-T2 MRI, illustrated by three different acquisitions performed on a phantom containing mineral oil with long T2 surrounded by pieces of rubber with short T2. All scans share the
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same nominal isotropic resolution, repetition time, and flip angle. (a) The gradient echo image is dominated by the oil signal while rubber is hardly visible. This is due to signal losses occurring before the echo is formed and acquired around the echo time TE. (b) In contrast, with zero echo time (ZTE)
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imaging, frequency encoding and acquisition start immediately after signal excitation, so that TE = 0. Hence, rubber appears in the image at full intensity. However, signal decay during the acquisition
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range Tk leads to image blurring, i.e., reduced actual spatial resolution for rubber. (c) Speeding up encoding by using a stronger gradient, thus creating a higher bandwidth BW, reduces the effects of signal decay and sharpens the edges of the rubber in the image. Hence, the two basic principles for
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encoding and acquisition in short-T2 MRI are 1) to start as early as possible and 2) to minimize the range Tk to avoid detrimental effects of signal decay. Fig. 4: High bandwidth as a requirement for short-T2 MRI. Imaging short-T2 samples at high spatial
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resolution requires both rapid encoding and large gradient (G) time-integrals. This can be achieved simultaneously only by using strong gradients, thus resulting in high resonance bandwidth. Fig. 5: Concepts and techniques for short-T2 MRI. The concepts follow the two principles of (1) an
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early start and (2) a small range of data acquisition, as illustrated in Fig. 3 and indicated by the red
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labels. (1) is achieved by A) using 3D imaging without slice selection, B) employing radial center-out gradient encoding, and C) performing signal excitation in the presence of the encoding gradient. For concept B, using either pure Cartesian phase or pure radial frequency encoding patterns favors one or the other principle. Combinations of the different excitation and encoding schemes lead to the four basic techniques for short-T2 MRI, namely constant time, single point, ultra-short TE, and zero TE
imaging, along with their intermediate relatives. CTI = constant time imaging, SPI = single point imaging, SPRITE = single-point ramped imaging with T1 enhancement, MP mapping = multipoint mapping [382], UTE = ultra-short echo time, BLAST = back-projection low angle shot, RUFIS = rotating ultra-fast imaging sequence, ZTE = zero echo time, WASPI = water- and fat-suppressed
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
proton projection imaging, SWIFT = sweep imaging with Fourier transformation, PETRA = pointwise encoding time reduction with radial acquisition, HYFI = hybrid filling, GM = gradient-modulated, RHE = ramped hybrid encoding. Fig. 6: Sequence diagrams of the four basic techniques for short-T2 MRI. They all employ 3D centerout k-space encoding from which one TR interval is depicted. (a) In constant time imaging (CTI), pure
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phase encoding is performed using a gradient of constant length and variable amplitude. One data
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point is acquired per excitation after identical TE, including k = 0. Thus, spatial resolution is not impaired by T2 decay but signal is reduced related to TE. (b) Shorter TE is achieved with single point
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imaging (SPI) where gradient ramps are excluded from encoding by performing both excitation and acquisition of high bandwidth while the gradient is at full strength. As indicated in the scheme by the
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grey bars, gradients need not be switched off between successive TR intervals, thus enabling nearsilent operation. (c) In contrast to the previous techniques, ultra-short echo time (UTE) imaging employs pure frequency encoding. Thus higher SNR efficiency is achieved but T2 decay may affect
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spatial resolution. After non-selective excitation, data acquisition is started concurrently with gradient ramping, the minimum TE being determined by the RF dead time. (d) In zero echo time (ZTE)
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imaging, high-bandwidth excitation is performed after ramping up the gradient. TE is actually zero because k = 0 coincides with the center of the RF pulse. Data acquisition starts only after the RF dead time, thus leading to a gap in central k-space. Equivalently to SPI, ZTE imaging can be performed
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with strongly reduced acoustic noise.
Fig. 7: Patterns in k-space of the four basic techniques for short-T2 MRI, schematically showing a central plane of 3D k-space for each case. The distance between neighboring k-space points according
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to the Nyquist criterion is given as the dwell time DW = 1/BW for the bandwidth BW in the FOV. (a) Both CTI and SPI are usually performed with Cartesian sampling. One k-space point is obtained per excitation after radial encoding of varying speed and direction. Thus, all points are equally T2-
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weighted but with a relatively long TE. (b) In UTE imaging, data are acquired along the radial
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trajectory. In central k-space (yellow), encoding speed increases while ramping up the gradient, leading to a non-uniform radial sampling density. (c) In ZTE imaging, full, constant encoding speed is achieved but data are missing in central k-space (red) due to finite RF dead time. Additional data points are acquired by radial oversampling. Fig. 8: Effect of gradient ramping on depiction behavior in UTE imaging. A 1D simulation is shown using different ramp durations associated with different gradient slew rates. All times and durations are normalized with respect to ' which is the acquisition range at constant gradient without ramping, i.e., as in the ZTE case shown in Fig. 4d. Using the parameters FOV 240 mm, spatial resolution 1 mm, and gradient strength 25 mT/m results in ' = 0.47 ms. A T2 value shorter than
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
' of 0.20 ms is assumed to emphasize the effects of signal decay. As a reference, the ideal case is shown where all data are obtained without any delay after signal generation, i.e., unaffected by T2. (a) Time of acquisition of the k-space positions of one radial readout. During the ramp, as indicated by the dashed lines, k-space speed increases until the target gradient strength is achieved. Compared to ', the UTE acquisition range is increased by half the ramping time. (b) Modulation transfer
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function (MTF) describing the apodization in k-space due to exponential T2 decay according to the
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delayed acquisition times shown in a). (c) Point-spread-function (PSF) obtained by symmetrically complementing, zero-filling, and Fourier-transforming the MTF in b). The displayed PSFs describe
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the distribution of signal from a pixel center (position 0) to other reconstructed pixels in the image and are depicted for the first 12 neighboring pixels. In the ideal case, the PSF is 1 at the location of origin
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and 0 at all other pixel positions. The Sinc shape reflects finite sampling up to kmax. For UTE encoding, the PSFs show a broadening of the main lobe and damped Sinc side lobes. (d) Normalizing
slew rates further emphasize this effect.
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the UTE PSFs to their maximum value reveals the associated loss of spatial resolution. Low gradient
Fig. 9: Impact of gradient fidelity on UTE image quality. (a) Ideal gradient shape with ramp. (b)
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Simulated image of a sphere obtained from data acquired with the ideal shape in a). (c) Distortion of gradient shape obtained for an exponential first-order, so-called gradient eddy current with time constant 0.05 '. (d) Image obtained when reconstructing data acquired with the distorted gradient
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in c) using the nominal, ideal trajectory. (e) Phase evolution from a zeroth-order field generated by an exponential, so-called B0 eddy current with time constant 0.5 '. (f) Image obtained from data with B0 eddy current in e) induced by the gradient coil encoding the horizontal direction. Correcting
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data for the B0-eddy-current-induced phase and reconstructing based on the k-space trajectory distorted by the gradient eddy current results in an image virtually identical to b). Fig. 10: Contributions to the initial RF dead time in pulsed MR. For symmetric RF pulses of constant
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frequency, the focal point is located in the pulse center, so that only half of its duration contributes to
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the dead time. During T/R switching, the MR signal is corrupted by spikes and transients. After these RF events, fully filtered and uncorrupted data are only available after the lead-in time of the digital filter, corresponding to half of its length in the case of a symmetric filter with finite impulse response. The dwell time DW corresponds to the bandwidth in the target FOV according to Nyquist (see Fig. 7), whereas DWFOV is smaller and associated with radial oversampling. Fig. 11: Illustration of 1D algebraic reconstruction (AR) of a ZTE image. (a) Simulated ZTE k-space data obtained for a box-shaped object, exhibiting a gap of 3 Nyquist dwell times on the radius and two-fold oversampling. (b) Image obtained by FT of the data after filling the gap with zeroes, leading to low-frequency artifacts. (c) Image obtained by AR which is free of artifacts. (d) Data in k-space
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
obtained as FT of the reconstructed ZTE image, demonstrating recovery of data in the gap along with elimination of the oversampling. Fig. 12: Noise behavior in ZTE imaging. (a) 1D image reconstructed from pure noise time-domain data. Bad conditioning at a gap size of 3 Nyquist dwell times DW = 1/BW leads to noise enhancement and correlation in the image domain, which manifests as a low-frequency image distortion. Another
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realization of the noise input data would lead to a different shape in the image. (b) Using as input the
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same noise plus the signal from Fig. 11a results in a superposition of the two components also in the image. (c) Reconstruction of a 3D image from noise projections with gap = 3 DW as a). Due to high
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inherent oversampling of 3D radial data in central k-space, noise enhancement is averaged out. (d) Only at larger gap = 4, further enhanced and correlated noise patterns are observed also in the 3D
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image.
Fig. 13: Background signal in ZTE imaging. Images of a human head were acquired at 7 T using a
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T/R birdcage RF coil. All kinds of signals were detected, including signals from very short-T2 materials of the RF coil and cables, the head support, and the patient table. (a) With a small dead-time
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gap and a FOV adapted to the head size, ordinary aliasing of radial data is manifested at the image center and borders. (b) A FOV twice as large contains most of the background signal and aliasing is largely eliminated. However, scan time is four-fold, and bandwidth BW and the gap size given in Nyquist dwell times DW = 1/BW are doubled. (c) For a slightly larger gap, worse conditioning
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amplifies the aliased out-of-band signal. (d) Choosing a larger FOV does not help in this case as the increased gap of 3 DW also leads to amplification of residual aliased signal.
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Fig. 14: RF excitation for ZTE imaging for a given signal bandwidth BW in the FOV. (a) Time course of short and long block-shaped pulses and a modulated pulse. Durations are given in Nyquist dwell times DW = 1/BW. The frequency sweep of the modulated pulse approximately covers BW. (b) Spectral amplitudes exhibiting high uniformity for the short block (= hard) pulse and strong variations
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for the longer block pulse. The modulated pulse achieves about three times the amplitude of the hard
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pulse for a comparable uniformity. (c) ZTE image of a sphere using the short block pulse, showing uniform excitation. (d) A longer block pulse of duration 2.5 DW results in strong variations in image intensity. (e) A pulse of equal duration but with amplitude and frequency modulation provides uniformity comparable to the short block pulse. Fig. 15: RF excitation using a long modulated pulse as done in the SWIFT technique. To avoid missing signals generated and encoded during earlier parts of the pulse, excitation and acquisition are performed quasi-simultaneously. This is achieved by periodically alternating between the two operation modes. Thus both excitation and acquisition are run at a reduced duty cycle which is further
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
decreased by switching dead times. To also fully encode signals created at the end of the pulse, the encoding gradient and acquisition need to be continued after excitation [107]. Fig. 16: Patterns in k-space of two techniques filling the dead-time gap in ZTE imaging. (a) In PETRA, the gap area is covered by additional Cartesian SPI acquisitions. (b) In WASPI, additional radial projections are acquired at reduced gradient strength to limit the gap size to half a Nyquist
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distance. To limit extra scan time, the angular density is chosen to meet the Nyquist criterion only at
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the interface between the two parts of the data.
Fig. 17: Image fidelity of ZTE imaging and related techniques, illustrated in analogy to Fig. 8. (a) In
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SPI, the time of acquisition is constant as in the ideal case but delayed according to the largest
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encoding. ZTE encoding starts at 0, increases linearly, and ends at the SPI delay. Note that with the present gap size of 50 μs = 13 DW, the ZTE case is not actually possible due to ill-conditioning and is only shown to illustrate the purely exponential decay. With PETRA, the additional SPI data in the gap
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are acquired after the dead time, leading to a reduced acquisition range Tk. In WASPI, a discontinuity is created by the different gradient strengths. (b) T2 decay illustrated by the MTF shows no effect for
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SPI, exponential weighting for ZTE acquisition, a flattened initial part for PETRA, and a discontinuous behavior for WASPI. (c) The PSF shows that at the original pixel location the strongest signal reduction is experienced by SPI, whereas ZTE, WASPI, and PETRA exhibit a comparable decrease. (d) The normalized PSF shows ideal resolution for SPI and an increase of the main lobe
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width for the three ZTE relatives. For both ZTE and PETRA, a resolution loss is observed from the distribution of signals to neighboring pixels. With WASPI, this effect is stronger and reaches further
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out, thus leading to side lobes.
Fig. 18: Generating contrast in ZTE imaging by magnetization preparation demonstrated on an excised lamb joint. (a) Non-prepared ZTE image showing primarily proton density and some T1 contrast. (b) In gradient echo imaging, signals from short-T2 tissues are reduced, in particular from
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cortical bone. (c) In ZTE imaging with fat suppression, both short- and long-T2 water signals remain.
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(d) Selectively suppressing long-T2 water in the ZTE sequence removes signals from cartilage and
muscle, and fat signal dominates the image. (e) Applying suppression of both fat and long-T2 water results in positive contrast for short-T2 tissues, showing primarily trabecular and cortical bone. Remaining bright fat signal (right) indicates imperfect suppression due to local off-resonance. (f) In the CT reference image, strong similarities with the MRI data in e) are observed. Reproduced with permission from Ref. [153]. Fig. 19: Off-resonance effects in radial imaging, demonstrated for the specific case of chemical shifts in the presence of both water and fat. (a) Simulated images with a pixel bandwidth (PBW) corresponding to the frequency difference between water and fat. Setting the demodulation frequency
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
f to the water resonance results in sharp edges of the water compartment and blurred depiction of the fat parts. At interfaces of the two compartments, destructive and constructive interference lead to edge enhancement (arrow). Demodulation at the fat frequency changes the order of low and high signal intensities at interfaces whereas at the intermediate frequency, edges are not enhanced but blurred. (b) In-vivo ZTE data of a human knee acquired at 7 T with PBW = 870 Hz, corresponding to 0.87 times
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the water-to-fat-shift. At different demodulation frequencies, the equivalent effects to the simulations
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in a) are observed, as indicated by arrows.
Fig. 20: Resolution enhancement by exponential T2 filtering in simulated radial MRI with center-out
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encoding. (a) Two ellipse-shaped objects without T2 filtering. For a relatively long T2 equal to the acquisition range Tk, the object is sharply depicted. For a significantly shorter T2, blurring is observed.
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(b) Moderate T2 filtering enhances resolution of the short-T2 object but leads to edge artifacts for the object with longer T2. (c) With full compensation of T2 decay, the short-T2 object is sharply depicted.
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However, strong artifacts are observed for the long-T2 object and noise is enhanced.
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Fig. 21: Selected short-T2 MRI applications. All images were obtained using the 3D ZTE technique. (a) MSK: Human knee acquired at 7 T with isotropic spatial resolution of 0.75 mm in 4 m 17 s. Reproduced with permission from Ref. [205]. (b) Pulmonary MRI: Mouse lung acquired at 4.7 T with a resolution of 0.31 mm in 4 m 30 s using respiratory gating. Reproduced with permission from Ref.
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[99]. (c) Dental MRI: Excised human tooth acquired at 11.7 T with a resolution 0.13 × 0.13 × 0.18
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mm3 in 21 m 18 s. Reproduced with permission from Ref. [310].
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Table 1: Principles for short-T2 MRI. = time of acquisition, TE = echo time, SNR = signal-tonoise ratio, = range of times of acquisition.
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Principle for data acquisition 1. Early start
2. Small range
Characterized by
Affects
!" , TE
SNR
Spatial resolution
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Table 2: Examples of short-T2 tissues and materials. Relaxation times were derived and selected from single- or multi-component analysis. An asterisk * indicates a T2* value as opposed to T2. Unless stated otherwise, the values associated with tissues refer to 1H signal in water molecules. B0 [T]
Origin
T2 or T2* [μs]
References
Cortical bone
0.47
Porcine, bovine, specimen
400
[347]
(bone water)
0.64
Human, specimen
400
[348]
1.5
Human
420 – 500 *
Human, specimen
401 *
Human
260 – 786 *
Human, specimen
237 – 576 *
[224, 233, 349]
Bovine, specimen
290 – 368 *
[175, 227, 350-352]
4.7
Human, specimen Human, specimen
7.0
O
O
290 *
[354]
431 *
[168]
290 *
[236]
Human, specimen
302 *
[349]
Human, specimen
540, 400 *
[243]
Porcine, specimen
597 *
[258]
Human
176 – 207 *
[252, 355]
Bovine, specimen
190 *
[356]
Human
160 *
[249]
Human
178 *
[256]
Lamb, specimen
189 *
[250]
Synthesized
448
[248]
7.0
Chicken, specimen
675
[248]
9.4
Lamb, specimen
119 *
[250]
Rabbit, specimen
90 *
[22]
0.47
Porcine, specimen
650 – 3800
[25]
1.5
Human
1200 – 2000 *
[150, 259]
3.0
Human
5200 – 6000
[161]
Human
600 – 2320 *
[31, 54, 173, 227, 260, 357, 358]
Human, specimen
1180 – 2180 *
Bovine, specimen
1280 *
Human
340
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1.5
PR
Bovine, specimen
3.0
R N U JO
[353, 354]
[349]
5.6
Tendons
400 – 416
368 *
11.7
(bone matrix, hydroxyapatite 31P)
[53, 166, 172, 231]
Human, specimen
9.4
Cortical bone
[349]
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Mouse
[232]
PR
3.0
F
Tissue/material
7.0
[359, 360] [350] [361]
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Collagen protein
Dentine
[51]
3.0
Human
5800 – 9300
[161]
Human
3260 *
[173]
Caprine, specimen
1530 *
[350]
1.5
Human
3000 *
[150]
3.0
Human
9700 – 11400
[161]
Human
4190 *
Human, specimen
1800 – 2100 *
1.5
Human
5000 – 10000
3.0
Human
1340 *
Human, specimen
480 – 990 *
9.4
Bovine, specimen
0.47
O [51]
O
[173]
[264, 350]
Porcine, bovine, specimen
12 *
[347]
0.7
Bovine, specimen
20 *
[363]
1.4
Rat tendon, specimen
90
[32]
4.7
Human, specimen
57, 12 *
[353]
9.4
Rabbit, specimen
20 – 40 *
[228]
Human, specimen
38000
[364]
Human, specimen
12, 65, 200, 1000 *
[364]
Human
324
[307]
Human, specimen
12, 65, 230 *
[365]
0.9
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0.9 1.5
Human
150 – 350 *
[366]
3.0
Human
358 *
[320]
Human
420 *
[147]
Human
500 – 700 *
[324]
Human
260 *
[367]
Human
336 *
[368]
Human, MS specimen
216 *
[320]
Ovine, specimen
200 – 300 *
[369]
Ovine, specimen
150 – 400 *
[370]
Bovine, specimen
114 *
[320]
4.7
Rat, specimen
50 – 1000 *
[318]
7.0
Human
200 – 300 *
[324]
9.4
Human, specimen
38 – 57
[371]
Rat, bovine, specimen
8 – 1000 *
[35]
U JO
[227, 350]
[362]
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Myelin lipids
[173]
2300
3.0 Dental enamel
F
4000 – 10000
PR
Cartilage
Human
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Meniscus
1.5
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Ligaments
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Human
4000 *
[372]
0.68
Rat, specimen
2340 *
[373]
1.5
Human
1100 – 2110 *
[287, 293, 374-377]
2.0
Rat, specimen
830 *
[373]
3.0
Human
460 – 820 *
[284, 293]
Mouse
910 – 1000 *
[288, 292]
4.7
Mouse
460 – 500 *
7.0
Mouse
395 *
7.9
Mouse
770 *
Liver
3.0
Human
300 – 450 *
Rubber
2.5
Polyvinylchloride (PVC)
2.4
Acrylonitrile butadiene styrene (ABS)
2.5
R N U JO
O
[162] [167]
O
[379, 380] [381]
180 – 1350 *
[147, 172, 223]
38 *
[340]
98 *
[340]
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2.4
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Polyethylene (PE)
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2.4
[285, 378]
300 – 1300 *
3.0 Polymethylmethacrylate (PMMA, acrylic glass)
F
0.5
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Lung
442 *
[340]
760 *
[381]
JOURNAL PRE-PROOF
a
Molecular tumbling or crystal rotation Dipolar field
B0 θ
F
θ
B0 + ΔB(θ)
Spectrum
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Liquid
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b
O
O
Molecule
Single crystal
θ = 20°
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Liquid with relaxation
θ = 90°
Powder
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θ = 55°
f
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0
θ = 0°
Free induction decay
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c
0
Fig. 1
t
Fig. 2
ΔB
Amplitude
Interaction (e.g. dipolar coupling)
τc
t [s]
τc = 0.05 ns
τc = 0.4 ns
W0
W2
AL W1
W 1 ~ J(ω0)
W 2 ~ J(2ω0)
Transition probabilities
W 0 ~ J(0)
FT
[s]
Liquids
F
O
τc [s]
τc = 1/ω0
Relaxation
ω [rad/s]
T2
T1
Spectral density
0 ω0 2ω0
J
O
PR
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Weights
PR τ [s]
Autocorrelation function
Energy diagram spin pair
AC
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U
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Random field
Solids
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Gradient echo
t
TE
AQ
Signal loss
Rubber T2 ≈ 0.4 ms
Oil T2 ≈ 30 ms
b ZTE
AQ
1 Early start
TE = 0
Tk
c ZTE
F
2 Small range
O
AQ
Tk
Tk = 0.32 ms, BW = 200 kHz
O
PR Signal decay
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Tk = 1.28 ms, BW = 50 kHz
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TE = 1.32 ms, BW = 100 kHz
Fig. 3
G
RF
Sig
a
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Fig. 4
+
High bandwidth
F
O
O
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Large G-integral
High resolution
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Strong G
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Rapid encoding
Short T2
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U
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Fig. 5
1
Radial center-out encoding
B
1
Frequency encoding Radial
Intermediate
2
Phase encoding Cartesian
C
A
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Ultra-short TE UTE
Constant Time Imaging CTI
GM-SWIFT
MP mapping WASPI PETRA HYFI
Single Point Imaging SPI, SPRITE
RF during G
3D
2
1
1
Zero TE BLAST, RUFIS, ZTE SWIFT
F
O
O
PR
RHE, GM-PETRA
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Intermediate
No slice selection
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RF before G
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Concept
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TE
AQ
Tk
PR
TE = 0
F
AQ
Tk
TE
O
O
Dead time
d ZTE
b SPI
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k=0
AQ
Tk = 0
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TE
U
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Ramp
c UTE
Fig. 6
G
RF
G
RF
a CTI
AQ
Tk = 0
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gradient encoding
b
k-space points:
DW
Nyquist
k0
F
O
missed
Dead time gap
ZTE
oversampled
O
c
PR acquired
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PR
UTE
Gradient ramp
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U
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CTI / SPI
Equal, long TE
Fig. 7
a
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Fig. 8
b
ramp
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a
Tk
d
F
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O
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plateau
c
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Fig. 9
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F
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O
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Fig. 10
RF pulse
F
DW
O
O
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T/R switching Filter lead-in
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Dead time
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DWov
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Fig. 11
d
a
Dead time gap
IFT
FT
c
b
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k-Space
F
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O
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Image
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c
Fig. 12
3D
1D
a
gap = 3
gap = 3
Noise
d
gap = 4
gap = 3
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Signal + noise
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F
O
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Fig. 13
FOV 530 mm BW 500 kHz DW 2 μs
FOV 265 mm BW 250 kHz DW 4 μs
Gap 2
b
Gap 1
a
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Gap 1.5
c
Dead time 6 μs
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Fig. 14
b
a
-1
-0.5
0
BW = FOV
0
0.5
f/BW
1 t/DW
d
c
2.5 DW
0.4 DW
e 2.5 DW
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O
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BW
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Fig. 15
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AQ
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RF AQ Dead interval interval time
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Fig. 16
a
SPI
b
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Lower gradient
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WASPI
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PETRA
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b
gap
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a
d
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Tk
c
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fat
Gradient echo
d
c
Water sup
f
CT
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e Fat + water sup
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cartilage
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Fat sup
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trab. cort. bone bone
ZTE
muscle
Fig. 18
b
a
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fat water
F
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f = ffat
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f = (fwater+ffat)/2
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PBW = fwater - ffat
Fig. 19
b
a
f = fwater
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b
c
F
T2,filter = T2,short
O
O
PR
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T2,filter = 2 T2,short
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T2,short = 0.2 Tk
No T2 filtering
T2,long = Tk
Fig. 20
a
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Fig. 21
a
patellar tendon
b
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F
O
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pulp
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caries lesion
c
bone
dentine
enamel
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JOURNAL PRE-PROOF Short-T2 MRI: Principles and recent advances Highlights Overarching concepts structure the variety of short-T2 MRI techniques Relaxation in a nutshell
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Detailed description of ZTE imaging and its variants
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Dedicated hardware requirements for all scanner components
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PR
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Comprehensive literature overview including T2 and T2* values of potential imaging targets
S0079-6565(19)30033-0 https://doi.org/10.1016/j.pnmrs.2019.07.001 JPNMRS 1483
To appear in:
Progress in Nuclear Magnetic Resonance Spectroscopy
Received Date: Accepted Date:
4 June 2019 26 July 2019
Please cite this article as: M. Weiger, K.P. Pruessmann, Short-T2 MRI: Principles and recent advances, Progress in Nuclear Magnetic Resonance Spectroscopy (2019), doi: https://doi.org/10.1016/j.pnmrs.2019.07.001
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© 2019 Elsevier B.V. All rights reserved.
JOURNAL PRE-PROOF Short-T2 MRI: Principles and recent advances Markus Weiger*,[email protected], Klaas P. Pruessmann
Institute for Biomedical Engineering, ETH Zurich and University of Zurich, Zurich, Switzerland *Corresponding
author at: Institute for Biomedical Engineering, ETH Zurich and University of
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O
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Zurich, Gloriastrasse 35, CH-8092 Zurich, Switzerland.
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Graphical abstract Highlights
Overarching concepts structure the variety of short-T2 MRI techniques
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Relaxation in a nutshell
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Detailed description of ZTE imaging and its variants
Dedicated hardware requirements for all scanner components
Abstract
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Comprehensive literature overview including T2 and T2* values of potential imaging targets
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Among current modalities of biomedical and diagnostic imaging, MRI stands out by virtue of its versatile contrast obtained without ionizing radiation. However, in various cases, e.g., water protons in tissues such as bone, tendon, and lung, MRI performance is limited by the rapid decay of resonance
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signals associated with short transverse relaxation times T2 or T2*. Efforts to address this shortcoming have led to a variety of specialized short-T2 techniques. Recent progress in this field expands the
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choice of methods and prompts fresh considerations with regard to instrumentation, data acquisition, and signal processing. In this review, the current status of short-T2 MRI is surveyed. In an attempt to
structure the growing range of techniques, the presentation highlights overarching concepts and basic methodological options. The most frequently used approaches are described in detail, including acquisition strategies, image reconstruction, hardware requirements, means of introducing contrast, sources of artifacts, limitations, and applications. Keywords: Relaxation; Zero echo time (ZTE); Ultra-short echo time (UTE); Single point imaging (SPI); Solid tissues
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
1 Introduction Magnetic resonance imaging (MRI) is a well-established and widely used modality in medical diagnostics due to its ability to depict tissues in three dimensions, with versatile contrast, and without ionizing radiation. It is based on spatial encoding and detection of macroscopic transverse nuclear magnetization. The latter is of limited lifetime and usually decays approximately exponentially with
F
time constants T2 or T2*, depending on the exact nature of the decay. Nuclei of molecules in liquids, in
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particular water protons in most soft tissues, are characterized by relatively long T2 in the range of tens of milliseconds to seconds, which readily permit extensive spatial encoding by gradient fields and
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full signal detection.
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However, much shorter signal lifetimes can be observed for nuclei in molecules with more restricted mobility. Some tissues, including tendons, ligaments, bone, and teeth, are stringy or solid and owe their short T2 to low water mobility. 1H and other nuclei bound in large molecules such as lipids and
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proteins can be even less mobile and thus exhibit yet shorter T2. However, loss of macroscopic transverse magnetization can also stem from dephasing due to B0 inhomogeneity at the microscopic
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scale. A special case is lung tissue, in which rapid decay of the water signal associated with strong, microscopic variation in magnetic susceptibility is further influenced by diffusion. Generally, the exact mechanism of coherence loss is relevant when considering methodology. However, in the context of imaging, in particular in vivo, many of the main considerations hold equally for different
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types of signal decay. Therefore, unless stated otherwise, all kinds of coherence loss (such as T2 or T2*) will be jointly referred to as T2 decay in the scope of this work.
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Rapid signal decay hampers MRI in two ways. In conventional scans, signals with T2 below approximately one millisecond are largely missed during the dead time between MR excitation and detection, causing voids in resulting images. In addition, even if short-T2 signal is still captured, signal decay during data acquisition translates into apodization in k-space, entailing loss of spatial resolution
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or even artifacts.
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To expand the scope of application of MRI to short-T2 tissues, continuing efforts have been made to
image such components by means of dedicated techniques. In fact, the field of short-T2 imaging is
almost as old as MRI itself. In solid-state NMR, short T2 is often addressed by line-narrowing techniques [1-4]. Unfortunately, line-narrowing is barely applicable in vivo since it involves either rapid sample rotation or high levels of radiofrequency (RF) power absorption. Therefore, short-T2 imaging in vivo must primarily rely on sufficient speed of spatial encoding and data acquisition. The chief traditional strategies to this end are single-point imaging (SPI) [5] and ultra-short echo time (UTE) imaging [6], which both employ efficient center-out k-space encoding. Over the past decade, the field has been expanded by zero echo time (ZTE) methods [7], of which an array of
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
implementations is now in use while more variants continue to emerge. Promising results of recent applications have promoted general interest in short-T2 imaging [8] and instrument manufacturers are increasingly following up with commercial implementations of new advances. At present, short-T2 MRI is a diverse field, and it can be a challenge to gain an overview and keep track of ongoing developments. Nevertheless, all of the techniques that are involved have emerged
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from a finite set of basic ideas that guide sequence design, signal detection, data processing, and
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hardware requirements. The goal of this article is to review the existing variety of short-T2 MRI methods from such a conceptual point of view, to work out commonalities and key differences
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between techniques, and to place them in a shared framework. It strives to offer guidance for getting started in short-T2 method development and for choosing the right approach towards a given
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application.
The presentation starts with a brief recapitulation of the physics underlying signal decay in MRI. It
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then discusses the basic principles of short-T2 imaging, which serve as a framework for a subsequent overview of existing techniques. Particular attention is given to two classes of techniques that are
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currently most employed in vivo, the UTE and ZTE approaches. A separate section is dedicated to specific hardware requirements. The final section of this review surveys the realm of current and potential applications.
2.1
Preamble
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2 Relaxation in a nutshell
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The general topic of relaxation is a complex branch of NMR physics. For the purpose of this article, no attempt is made to survey it fully or at significant depth. The aim of this section is rather to set the scene for MRI with short-lived signals by a brief, simplified review of the chief causes of rapid signal
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decay. For a thorough treatment of the topic we refer to NMR textbooks [9] as well as original work
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[10-12] and pertinent reviews [13-16]. 2.2
Overview
Relaxation is the process that drives a perturbed spin system back to thermal equilibrium with zero transverse
magnetization,
,
and
finite
longitudinal
magnetization,
.
A
phenomenological, mono-exponential description of relaxation is given by the Bloch equations. In the absence of an RF field they reduce to
(1)
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
(2)
with relaxation times T1 and T2, which vary greatly across different nuclei and their environments in given tissues and materials. The purpose of relaxation theory is to describe the mechanisms underlying relaxation and to relate T1 and T2 to more fundamental physical properties.
F
In most practical cases, relaxation of nuclear spins is based on two key ingredients: 1) some magnetic
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interaction of the spins in question with their surroundings that depends on orientation or position of the molecules containing the spins and 2) quasi-random molecular tumbling. Tumbling motion
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translates interaction into magnetic field fluctuations at the spin locations. These fluctuations cause continuous slight variations in the precession frequency of the spins and in their angle with respect to
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the static field B0, leading to irreversible loss of transverse magnetization and recovery of longitudinal magnetization, respectively. Interactions
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2.3
Relevant interactions occur within molecules as well as between spins and their wider environment.
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To contribute to relaxation, interactions must be spatially anisotropic or inhomogeneous. In particular, intra-molecular interactions must depend on the orientation of the molecule with respect to the B0 field whereas inter-molecular interactions must change with distance or relative orientation. The
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usually strongest and therefore most important interactions are: 1) Dipolar coupling: The dipolar field associated with a spin’s magnetic moment contributes to the local field at the sites of nearby spins. Dipolar coupling is a through-space effect that drops rapidly
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with distance and therefore acts mostly intra-molecularly. It will be discussed in more detail below. 2) Chemical shift anisotropy: In electron shells of molecules, the external B0 field induces currents, which in turn add to the total field experienced by nuclear spins. This effect is generally anisotropic
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because the topology of a molecule favors certain electron paths. The strength and direction of the
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created field thus depend on the orientation of the molecule relative to B0. With respect to relaxation, this effect is most relevant in solids. In liquids and gases, it is typically canceled by motional averaging. 3) Electric quadrupole coupling: In nuclei with spin > 1/2, the distribution of charge is non-isotropic,
giving rise to an electric quadrupole moment that interacts with electric field gradients generated by surrounding electrons. This interaction depends on the relative orientation of the molecule and the nuclear moments and thus fluctuates upon tumbling in the B0 field. The effects of orientation-dependent interactions are schematically illustrated in Fig. 1 for the case of dipolar coupling between two equivalent nuclear spins in a molecule. Upon rotation of the molecule,
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
the spins remain aligned with B0 such that their dipolar interaction changes (Fig. 1a). The associated effect on NMR spectra depends on the configuration and mobility of molecules in a sample (Fig. 1b). For a rigid single crystal, all local lattices form the same angle with B0, leading to a single sharp peak. A second symmetric peak is created by opposite spin alignment. In a powder, all lattice orientations co-exist and give rise to a continuum of resonance frequencies forming the characteristic Pake pattern
F
of a powder spectrum [17, 18]. In liquids, tumbling of molecules partly cancels dipolar interaction by
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temporal averaging. In the limit of rapid tumbling, the spectrum converges to a single sharp peak as would be obtained without any coupling. However, in real liquids, tumbling occurs on a finite time
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scale, still averaging dipolar coupling and leading to a single spectral peak, yet with a finite linewidth (Fig. 1b), which is associated with relaxation as discussed in the following section.
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Free induction decay (FID) signals corresponding to these spectra are shown in Fig. 1c. The hypothetical sharp peak corresponds to a signal of constant amplitude. With relaxation, the signal decays exponentially. As this decay is driven by stochastic fields, it is irreversible and cannot be
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countered by spin-echo refocusing. The two sharp peaks of the single crystal correspond to oscillation with a constant envelope in the time domain while the broad, sharply delineated spectrum of the
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powder entails a rapid signal drop with subsequent ringing. In the ideal powder, the loss of coherence is not of stochastic nature. In principle, it could be fully recovered by refocusing and is thus not considered actual relaxation. Mechanism of relaxation
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2.4
The dependence of relaxation on random molecular motion is shown in the final plot of Fig. 2
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(“Relaxation”, bottom right): In liquids, high molecular mobility leads to long T2 and T1 whereas in solids restricted mobility causes short T2 associated with long T1. The whole Fig. 2 summarizes the mechanism of relaxation driven by orientation- or position-dependent interaction in combination with
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tumbling motion. Quasi-random rotational or translational motion and the anisotropic or non-uniform field associated with interaction lead to a random field ΔB(t) at the location of a considered spin. The
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amplitude of the field fluctuations is given by the interaction whereas the temporal behavior is linked with the speed of the tumbling motion and characterized by the correlation time τc. The latter can be
determined by calculating the autocorrelation function of the random field which is usually well described by an exponential function. The Fourier transform (FT) of the autocorrelation function is the spectral density J(ω) of the tumbling motion, i.e., the energy available as a function of frequency. It has the shape of a Lorentz curve and depends on τc as illustrated by two examples. The spectral density is the key link between molecular motion and relaxation. Relaxation changes the state of the spin system. Longitudinal relaxation affecting Mz is associated with change of the populations of the different energy levels, as indicated for a two-spin system in
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Fig. 2. The transition probabilities depend on the order of the state change and are proportional to the spectral density of the tumbling motion at multiples of the Larmor frequency ω0. 2.5
T1 and T2
The final plot in Fig. 2 shows the qualitative behavior of relaxation times as functions of the
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correlation time of tumbling motion. At very short τc as found in small-molecule liquids, interactions are averaged effectively, resulting in very slow relaxation, i.e., long T1 and T2. As τc increases, both
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types of relaxation initially speed up. T1 reaches a minimum at τc ≈ 1/ω0, reflecting dependence on
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J(ω0) and J(2ω0). Hence, longitudinal relaxation is most effective when 1/τc approximately matches the Larmor frequency, similar to the resonance condition for RF excitation. It is instructive to view
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longitudinal relaxation as caused by the transverse components of the random field whose action on Mz starts to cancel itself once spins perform more than half a precession cycle during the correlation time [14]. In contrast, T2 has no minimum, reflecting the fact that for large τc it is dominated by J(0).
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In this regime, transverse relaxation is driven mostly by the longitudinal component of the random field. Longitudinal relaxation inherently entails a certain loss also of transverse coherence, limiting
2.6
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the transverse relaxation time to T2 ≤ 2T1. In most practical cases T2 ≤ T1 [19]. T2*
In real experiments, the decay of an FID signal obtained from a given volume is often faster than
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according to T2. The chief cause is spatial variation of the static magnetic within the volume. The associated frequency spread causes dephasing of the spins, i.e., loss of macroscopic transverse coherence and thus signal decay. As long as the field variations are static, the dephasing can, in
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principle, be undone by spin-echo refocusing. In contrast to pure random-field-driven relaxation, signal decay including macroscopic dephasing has been termed apparent transverse relaxation. The shape of such decay varies widely depending on the underlying field distribution. Nevertheless, it is
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frequently approximated by an exponential with the reduced time constant T2*. In the following, three
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cases of apparent transverse relaxation are discussed more closely: 1) Dipolar coupling in a powder as described in Fig. 1. In this case, the shape of the FID deviates strongly from exponential behavior. Large frequency spreads on the order of typically tens to hundreds of kHz require large-bandwidth, high-power RF pulses for refocusing. Alternatively, the dipolar interaction may be averaged out by spinning the sample at the magic angle or eliminated by RF decoupling pulses [18]. Note that such approaches are usually impractical in vivo. 2) Microscopic structure: Biological tissues (e.g. trabecular bone) and other samples of interest (e.g. porous rock) often exhibit microscopic structure on a length scale below feasible MRI
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
resolution. Related microscopic variation of magnetic susceptibility causes corresponding field inhomogeneity and signal dephasing. In T2*-weighted MRI, the associated signal decay alters image intensity on a per-pixel basis and can be used to characterize the underlying structure and materials involved. 3) Macroscopic B0 inhomogeneity: Macroscopic compartments filled with materials of differing
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bulk susceptibility such as soft tissue, bone, and air, give rise to long-range field variation
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that also causes dephasing and signal decay. However, such dephasing is not characteristic of the tissue in which it occurs. For compartment sizes in the order of or larger than the spatial
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resolution, macroscopic susceptibility effects rather hamper MRI by depiction errors, e.g., signal shifts or blurring, and unspecific signal dropout. These problems can be mitigated by
2.7
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increasing spatial resolution and the strength of gradient fields used for spatial encoding. Diffusion
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In the presence of static microscopic field inhomogeneity, signal decay is further influenced by diffusion of the molecules that bear the resonant spins. Random motion translates field inhomogeneity
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into random field fluctuation that is individual for each spin and thus affects the nature of dephasing among the spin ensemble. This mechanism differs from relaxation as discussed above in that the effective field fluctuations are smaller by many orders of magnitude and arise from fields on a greater length scale such that their correlation times are much longer. As a result, dephasing under diffusion
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is amenable, at least in part, to refocusing by radiofrequency pulses. In a linear field gradient with amplitude G and unbounded Gaussian diffusion, the signal amplitude of an FID decays according to
(3)
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with self-diffusion constant D and gyromagnetic ratio γ [20, 21]. In spin-echo experiments with echo
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time TE, refocusing partly recovers the decay, resulting in an apparent relaxation rate of
(4)
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Hence, MR experiments may be hampered by increased apparent relaxation due to diffusion but can also give access to the physical property D. 2.8
Examples
As discussed above, various mechanisms can cause loss of coherence and thus rapid signal decay in tissues and materials. Accordingly, short-T2 MRI has a large variety of potential imaging targets, including the following prominent examples that will be further discussed in Section 10:
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
1) Solids: In solids, molecular motion is strongly restricted, leading to long correlations times. As seen in the relaxation plot of Fig. 2, in such substances short T2 is associated with long T1. The latter has two key implications for sequences with rapidly repeated RF excitations as introduced later: 1) steady-state magnetization is generally low and 2) maximum steady-state signal is reached with small excitation flip angles. In the biomedical context, a solid of
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particular interest is hydroxyapatite in the bone matrix [22].
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2) Bound water: Water molecules are small, leading to short correlation times and slow relaxation in free water. However, when bound to macromolecules or solids, water molecules
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exhibit reduced mobility and much shorter T2. An example is collagen-bound water in bone [23].
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3) Water in collagen fibers: Some tissues (e.g. tendons, cartilage) are composed of collagen fibers which are organized in directed, parallel bundles. Motion of water embedded in such highly ordered structures is strongly restricted, leading to short T2. In addition, there are
E-
preferred orientations along which bound water molecules line up. Thus the effect of dipolar coupling and hence T2 become a function of the angle between fibers and B0 [24-33]. In T2-
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weighted acquisitions, signal varies and peaks at the so-called magic angle of 55°. This needs to be considered for image interpretation but also holds potential for contrast manipulation. 4) Myelin: This membrane tissue, which plays a key role in the nervous system (see section 10), is largely composed of phospholipids arranged in liquid crystalline bilayers. Intra-molecular
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dipolar interactions cause rapid T2 decay of 1H signal, which varies with the orientation of the bilayer with respect to B0. As in typically probed volumes all orientations exist, the resulting
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signal is well described by a superposition of exponential decays with a wide range of T2 values, leading to so-called “Super-Lorentzian” lineshapes [34, 35].
5) Lung parenchyma: Signal decay of water in pulmonary tissue is an important example for the diffusion-driven mechanism described in section 2.7. Microscopic tissue-air interfaces at the
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level of the alveoli create local field gradients that rapidly dephase spins in water diffusing in
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the tissue [36]. This results in fast signal decay in gradient-echo techniques, which can be partly reversed with spin-echo sequences, according to Eqs. 3 and 4.
3 Fundamentals of short-T2 MRI
In this Section, the specific requirements for depicting tissues with short T2 are expressed in terms of two guiding principles. Starting from these principles, concepts of sequence design are derived, leading to four basic templates of short-T2 techniques.
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3.1
Principles
Figure 3 illustrates the challenges of short-T2 imaging. Common echo sequences hardly capture any signal from short-T2 spins due to signal loss by a factor
until the echo is formed and acquired
around the echo time TE (Fig. 3a). For gradient echoes this holds equally for pure T2 decay and T2*
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decay. In spin echoes, the latter is partly refocused, which can be beneficial when T2* is much shorter than T2. However, at very short T2*, RF refocusing requires commensurately large bandwidth and thus
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power deposition, which limits the applicability of spin echoes for this purpose. Therefore this option
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is not further considered here in favor of techniques that can equally handle short T2* and T2 . To contain the signal loss observed with echo techniques, dedicated short-T2 sequences must start
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gradient encoding and data acquisition quickly after signal excitation. Following this first principle renders short-T2 materials visible to begin with (Fig. 3b). However, signal decay during data acquisition still impairs spatial resolution. In mathematical terms, this is due to apodization in k-space,
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which is equivalent to convolution with a kernel of reciprocal width in the image domain [37]. To avoid such resolution loss, all image data must be acquired within a small time range after excitation.
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In Fig. 3c, this second principle is implemented by use of a stronger gradient, which shortens the readout and thus improves the sharpness of short-T2 features. More formally, the relationship between time and k-space in an MRI sequence is described by
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defining the time of acquisition as the time after signal excitation at which a k-space point is acquired. In echo-based MRI techniques, the value of this function at the k-space center, , is the echo time TE. The same definition of TE has widely been adopted for short-T2 imaging even though
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most of the techniques work without actual echo formation (see Section Error! Reference source not found. for details). To keep with this convention, the same notion is used in the present work. The range of times across which data are collected shall be referred to as the acquisition range
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!" .
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Based on these definitions, the principles of short-T2 sequences can be narrowed down to 1) data acquisition must start early, i.e., !" must be small, and 2) the acquisition range must be
small. Both of these conditions must hold on the time scale of relevant T2. Table 1 summarizes these considerations along with characteristic sequence parameters and their impact on image quality. Note that !" in short-T2 imaging is often, but not necessarily, equal to TE. To approach the encoded, nominal resolution, the acquisition range should approximately match the decay time constant, i.e., # [38]. Shortening the acquisition further will yield diminishing returns in terms of resolution but comes at the expense of signal-to-noise ratio (SNR).
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The principles have an immediate consequence on any short-T2 sequence, as illustrated in Fig. 4. Competing with fast signal decay by reducing TE and/or requires rapid signal encoding. On the other hand, achieving high spatial resolution requires large time integrals of gradient amplitude. These demands can be fulfilled jointly only by using strong gradients and corresponding large spread of resonance frequency across the imaged object. As discussed in the following, such strong gradients
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may be active during RF excitation and/or data acquisition, making large bandwidth a characteristic of
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short-T2 MRI. This consideration also reveals the fundamental challenge that necessarily short observation times of rapidly decaying transverse magnetization are equivalent to high-bandwidth
Concepts
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3.2
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acquisition. Hence, short-T2 imaging is generally more SNR-limited than conventional MRI.
At the level of sequence design, the principles discussed above lead to three chief concepts that recur across the variety of short-T2 techniques:
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A) No slice selection: Slice selection under a gradient of finite strength requires RF pulses of suitably small bandwidth and thus long duration, followed by slice rephasing. To avoid the associated delays,
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short-T2 imaging frequently works with non-selective volume excitation, omitting the slice selection gradient. All spatial encoding is then achieved by three-dimensional (3D) Fourier encoding. (An exception to this approach will be described later in Section 4.3.)
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B) Radial center-out encoding: To reach every targeted k-space point in the shortest possible time, kspace is traversed by straight radial trajectories that point outward from the center. This approach is
•
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used in two principally different ways:
Pure phase encoding: after each excitation, a single point in k-space is targeted and sampled once reached.
Pure frequency encoding: data acquisition is performed continuously along radial trajectories.
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•
C) Excitation under encoding gradient: To avoid losing time ramping up the readout gradient after
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excitation, the gradient can be switched on first such that it is already present at when excitation is performed. Figure 5 visualizes the combinations of these design elements and indicates published implementations using the respective approaches. While all techniques share concepts A and B, they differ with respect to the options for concepts B and C. All current short-T2 techniques can be represented in this scheme and thus viewed in a shared conceptual framework. As the figure indicates, two options each with respect to concepts B and C give rise to four basic sequence templates, which will now be considered in more detail.
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3.3
Basic sequences
The four basic schemes of short-T2 sequences and their associated k-space sampling patterns are shown in Figs. 6 and 7, respectively. All techniques employ non-selective hard-pulse excitation followed by 3D radial center-out encoding in k-space. Pure phase encoding is used in constant time imaging (CTI) [39] (Fig. 6a) and SPI [5, 40, 41] (Fig.
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6b). In both, a gradient of varying amplitude and direction is used to travel to one point in k-space per
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excitation (Fig. 7a). Each k-space point is sampled at the same fixed = TE, which must be chosen large enough to permit reaching the maximum k-space excursion that corresponds to the targeted
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resolution. While k-space is traversed radially, the points sampled are typically chosen to lie on a
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Cartesian grid. By virtue of fixed , all data points have identical T2 weighting. Therefore, signal decay prior to TE does not impair spatial resolution. However, it does affect the SNR in resulting images, most so for short-T2 material. In CTI, TE includes time for ramping the encoding gradient up
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and down such that it is off during excitation and acquisition. This has the advantage that the bandwidth of both can then be limited to the intrinsic frequency spread in the object. However, it
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exacerbates the SNR penalty of signal decay.
As a remedy, in SPI TE is reduced by performing RF excitation and data acquisition in the presence of the gradient (Fig. 6b). The bandwidth of both must then be increased by making them suitably short. As a result, the SNR yield of SPI is subject to competing influences of signal decay and acquisition
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bandwidth. It is optimal at , requiring a maximum gradient amplitude of $ %&
for given T2 and spatial resolution %& [40]. In addition to its short-T2 capabilities, the SPI approach
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offers relatively silent operation if gradients are not switched off between excitations but merely adjusted to a nearby value [42]. The key advantage of purely phase-encoded techniques is full nominal spatial resolution at short T2.
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Their chief drawbacks are long scan times and low SNR efficiency associated with point-wise sampling, low acquisition duty cycle, and comparatively long TE. For these reasons, pure phase
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encoding is rarely used in vivo [43]. It is mainly applied for imaging materials [44, 45] and, in MR microscopy, to mitigate diffusion effects [46]. Recently, as discussed in Section 6, partial SPI acquisition has found additional use as a means of complementing other, more time-efficient techniques. The terminology of purely phase-encoded imaging is not entirely consistent, with somewhat variable use of the terms CTI and SPI in the literature. With pure frequency encoding, lines in k-space are sampled by continuous signal acquisition during radial encoding. In this fashion, shorter TE, faster overall scanning, and higher SNR efficiency are accomplished. For an image matrix size M, the gain in acquisition duty cycle relative to SPI leads to
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an increase in SNR efficiency of at equal maximum G and of at optimized G (see 9.2). In both UTE [6, 47-54] (Fig. 6c) and ZTE [7, 55-58] (Fig. 6d) imaging, radial encoding is performed by playing out gradients of fixed amplitude but direction varying from repetition to repetition. Thus one spoke is obtained per excitation (Figs. 7b and 7c). In the UTE sequence, data acquisition is started concurrently with ramping up the gradient shortly after excitation. The minimum TE then amounts to
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half the RF pulse duration plus the dead time required to change from transmit to receive operation.
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Due to gradient ramping, the central k-space is traversed at lower, linearly increasing speed (Fig. 7b). This is avoided in the ZTE sequence where RF excitation is performed only after switching on the
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gradient, thus achieving instant, full encoding speed. As a consequence, similar to SPI, RF excitation must be broadband and near-silent operation is possible by only adjusting and not switching off the
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gradient between excitations. In the ZTE case, spatial encoding starts immediately at the time of creating transverse magnetization; hence the time TE at which it leaves the k-space center is effectively zero. However, during the finite RF pulse duration and the RF dead time initial signal is
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lost, leading to a gap in central k-space (Fig. 7c), which needs to be addressed by additional measures (see 5.3 and 6.1).
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For in-vivo MRI, UTE and ZTE imaging and their relatives are currently the most suitable and most widely used techniques and will therefore be discussed in more detail in Sections 4 and 5. For all implementations discussed above, the concept of radial encoding is central to short-T2
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depiction performance. With common Cartesian encoding, the principles stated initially cannot be implemented as effectively. Notwithstanding this, Cartesian gradient echo imaging methods can also
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be optimized for short-T2 imaging [59, 60]. 4 Ultra-short echo time imaging The UTE strategy is the most established and most widely deployed short-T2 approach in use today.
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While arguably simple in principle, UTE sequence design and image reconstruction deviate significantly from standard MRI. To obtain robust results, implementations must consider a number of
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subtleties and pitfalls as detailed below. 4.1
Gradient ramping
The basic idea of the UTE sequence as shown in Fig. 6c is to start data acquisition very soon after RF excitation while concurrently ramping up the radial readout gradient to its target amplitude. The gradient ramp plays a key role in UTE MRI because it affects image quality in several ways. Depending on the desired gradient strength and the available slew rate (see Section 9.2), the ramping part can engage a considerable portion of the total encoding time. For a given nominal spatial resolution, the encoding time must be extended by half the ramp duration relative to the ZTE case
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with full gradient amplitude throughout. The effect of the gradient shape on depiction behavior in UTE imaging is illustrated in Fig. 8, using the point spread function (PSF) as a criterion. The time of acquisition of each k-space value (Fig. 8a) shows sections of increasing and constant k-space speed during gradient ramp and plateau, respectively. Longer ramps lead to larger acquisition range . In conjunction with T2 decay, the acquisition timing translates into the modulation transfer function
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(MTF) (Fig. 8b), which is then Fourier-transformed to obtain the PSF (Fig. 8c). As ramp times
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increases, apodization of the MTF causes progressive broadening of the main lobe of the PSF. The
4.2
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normalized display (Fig. 8d) illustrates the associated loss of spatial resolution. Gradient fidelity
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Gradient ramping is also critical to UTE image quality in that the demanded gradient waveforms may not be produced with the necessary fidelity. Delays, finite gradient amplifier bandwidth, and eddy currents cause deviations from nominal gradient dynamics. These may include dynamic fields
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different from first spatial order, usually with a dominant zeroth-order component. The simulations in Fig. 9 illustrate effects of eddy currents on UTE image quality. The distorted gradient shape in Fig. 9c
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leads to severe alterations in signal intensity when image reconstruction is based on the nominal trajectory (Fig. 9d). So-called B0 eddy currents that generate zeroth-order field and thus spatially constant phase errors (Fig. 9e) result in edge artifacts (Fig. 9f). Both types of error are amenable to retrospective correction. Phase perturbation by B0 eddy currents can be removed from the acquired
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raw data while first-order field effects can be addressed by image reconstruction based on the effective, distorted k-space trajectory. In this way, a virtually artifact-free image is obtained,
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comparable to imaging with ideal gradient behavior (Fig. 9b). Models of system imperfection for retrospective correction may be set up at different levels of complexity ranging from mere gradient chain delays to general first-order field perturbation,
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additional zeroth-order [61], and second- and higher-order field terms [62]. Various strategies have been described for determining the parameters of such models, including approaches based on raw
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UTE data itself [63, 64], calibration scans [65-67], impulse-response characterization of the gradient system [68], and measurement of field evolution using modified MRI sequences [69, 70] or dedicated field probes [71] (see Section 9.5). 4.3
RF excitation
In the 3D UTE sequence shown in Fig. 6c, spatially non-selective RF excitation is followed by 3D radial encoding of the whole object. Alternatively, unlike SPI and ZTE techniques, UTE permits slice excitation under a separate selection gradient, forming a two-dimensional variant. To avoid increasing TE by longer RF pulses and slice refocusing, it has been proposed to employ half-excitation pulses
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with opposite slice gradients in two successive acquisitions [47, 72, 73]. Full slice selectivity is then obtained by summation of acquired data. On the downside, besides doubling the scan time this approach has been reported to increase sensitivity to eddy currents [74, 75] and motion. Relaxation during excitation must be considered when pulse durations are significant relative to T2 of interest. In-pulse relaxation reduces the actual flip angle and renders it T2-dependent [52]. The
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presence of an RF field changes the physics of relaxation, which must be taken into account in
Echo time
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4.4
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sequence optimization and the interpretation of data [76, 77].
Two different definitions of TE are common in the UTE literature. The first, as shown in Fig. 6c and
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used throughout this article, assumes that relaxation starts at the magnetic center or focus point of the RF pulse [77, 78]. For symmetric waveforms, including the commonly employed hard pulses, this is the center of the pulse. The second definition assumes that relaxation starts only after the RF pulse
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[38]. Both pictures are incomplete in that they do not account for relaxation during excitation as discussed in the preceding section. Hence, the full characterization of a short-T2 sequence also
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requires RF shape and duration in addition to TE of either definition.
In short-T2 samples, the SNR that can be achieved with UTE imaging depends on the shortest possible TE, which is governed by sequence implementation and hardware constraints (Tab. 1). A lower bound
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for TE is given by the RF dead time, which is composed of the fraction of the RF pulse duration after its magnetic center, the time required to switch between RF transmission and reception, and the leadin time of bandpass filtering [79] (see 9.3). On current MR systems, typical TE values for 3D UTE
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MRI are in the order of several tens of microseconds. In the UTE scheme, TE can also be varied, which is a useful feature. It permits creating T2* contrast, emphasizing short-T2 tissues by image subtraction [80], and mapping T2* quantitatively [81].
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Deliberate choice of somewhat longer than minimal TE can serve to avoid detection of signals that
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decay too fast for meaningful spatial encoding even with the UTE approach. Disturbing signal of such exceedingly short T2 most commonly originates from hardware parts and causes cloudy image background when unaddressed [82] (see 9.3). 4.5
Image reconstruction
Due to non-Cartesian sampling, image reconstruction from UTE data involves interpolation in kspace, which is usually performed with gridding procedures [83, 84]. In particular, this includes correction for non-uniform k-space density which can be accomplished by using a pre-calculated density correction function (DCF) [85-89]. One approach to obtain a DCF defines local k-space
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density as the area surrounding a sampled point as calculated by Voronoi diagrams [86]. For radial sampling, this approach leads to the so-called rho filter [90] for which an implementation is provided in Section 12.1. Density correction is an approximation based on splitting up reconstruction into FT and a preceding filter, the DCF. More rigorously, reconstruction is viewed as a general inverse problem. It can be
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solved efficiently using iterative algorithms [91] which inherently perform correction for non-uniform
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k-space sampling. In this case, application of a DCF can serve for pre-conditioning, thus accelerating convergence. The choice of the DCF then affects only convergence but not the accuracy of the
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reconstruction result.
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5 Zero echo time imaging
Image acquisition according to the ZTE scheme shown in Fig. 6d dates back several decades. Renewed recent attention has been driven by technical and methodological advances as well as
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increasing interest in short-T2 capability and silent MRI. To express its relation to the more established
5.1
Sequence properties
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UTE method, the approach has been dubbed ZTE imaging [7].
The basic ZTE sequence is an imaging analogue of the pulse-acquire NMR experiment. The only modification consists in spatial spreading of resonance frequencies by addition of a projection
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gradient that is present during RF excitation and FID acquisition. This approach was in fact used in the very first MRI experiment ever reported [92]. Only later it was identified and exploited as an
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approach particularly suitable for short-T2 imaging [55-58]. Relative to UTE and standard present-day MRI the distinguishing feature of ZTE imaging is RF excitation in the presence of the encoding gradient. While arguably a small difference in a sequence
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diagram, it is of great consequence with regard to sequence considerations, hardware requirements, data acquisition, and image reconstruction. The chief objective of excitation under the encoding
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gradient is to remove the need for gradient ramping while short-T2 magnetization evolves. It saves
time, offers immediate full k-space speed, and minimizes gradient switching even after data acquisition. In this way, enhanced short-T2 sensitivity is obtained in combination with reduced
acoustic noise or even virtually silent operation and robustness against eddy currents. Removal of gradient ramps also reduces the time required per repetition and increases the acquisition duty cycle. In combination with large gradient strengths as required for short T2, very short repetition time (TR) can be achieved and scan times for ZTE imaging are comparatively short [93].
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These capabilities come at the expense of shifted and partly increased hardware requirements. Continuous gradient operation for short-T2 imaging requires high gradient strength at full duty cycle. This demand is not commonly met by present-day gradient systems and has prompted developments in dedicated gradient hardware (see Section 9.2). On the other hand, the ZTE mode comes with reduced specifications in terms of gradient switching speed. In turn, it partly shifts hardware
Dead time
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5.2
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transmit-receive (T/R) switching are essential as will be detailed in the following.
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challenges from the gradients to the RF subsystem. In particular, high-bandwidth excitation and rapid
In pulsed MR experiments, signal excitation and acquisition are separated in time to accommodate the
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large difference in the RF amplitudes involved. Changing between the respective modes of operation takes time, resulting in a dead time after RF excitation during which MR signal is not yet detected. In many MR experiments, especially those involving echo formation, acquiring data immediately after
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RF excitation is not of interest. However, in short-T2 imaging, encoding and acquiring fresh MR signal as soon as possible is one of the key principles (see 3.1). With immediate encoding at full k-
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space speed, the ZTE sequence is arguably the most rigorous implementation of this rule and thus particularly vulnerable to data loss during the dead time. In 3D k-space it causes a spherical gap at the center (Fig. 7c). To contain the gap size one must minimize the dead time, which has contributions from the RF pulse, T/R switching, and digital filtering (Fig. 10) [7]. The chief means of limiting the
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gap are thus the use of short RF pulses [93] (see Section 5.5), speeding up T/R switching (see Section 9.3), and the design and use of short digital filters (see Section 9.4). The residual k-space gap must be
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addressed either by means of image reconstruction (see Section 5.3), by acquisition of additional data (see Section 6.1), or a combination of the two. Ideally, RF dead time would be avoided altogether by performing signal acquisition also during
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excitation. However, due to the large dynamic range between transmit and receive signal this
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approach is of limited practicality (see Section 9.3). 5.3
Algebraic Reconstruction
Straightforward image reconstruction from data with a gap in central k-space, using standard FT, leads to prohibitive artifacts of low spatial frequency. However, for small gap sizes the missing data can be recovered by taking two measures: 1) oversampling in the direct, radial dimension (Fig. 7c, Fig. 10) and 2) treating reconstruction as a general inverse problem set in the framework of linear algebra [94]. Standard sampling according to the Nyquist-Shannon theorem is associated with harmonic encoding functions that are mutually orthogonal within the field-of-view (FOV). In contrast, oversampling is associated with encoding functions that exhibit non-vanishing scalar products with
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the Nyquist set. In the algebraic framework, these dependencies permit the implicit formation of missing low-frequency harmonics by linear combination of available encodings [95]. This approach is known as finite-support or band-limited extrapolation and was previously used for low-frequency restoration in MRI in a different context [96]. It is limited to a relatively small range of extrapolation, beyond which it becomes impractical due to ill-conditioning (see Section 5.4). Greater oversampling
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improves conditioning somewhat but also increases the amount of resulting data. Therefore,
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oversampling factors range between 2 and 4 in practice [95].
Formally, the algebraic approach models data acquisition by means of an encoding matrix [97]. Its
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rows reflect the encoding functions actually implemented. Hence, the gap is represented by the absence of corresponding rows while oversampling elsewhere leads to an increased number of rows
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relative to standard Fourier encoding [7]. Finite support is implied by limiting the length of the rows to the FOV. Algebraic reconstruction is achieved by solving the encoding equation for the unknown image, eliminating the dead time gap and returning to regular sampling density in k-space. Figure 11
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illustrates this approach for the case of one-dimensional (1D) ZTE data. In principle, it can be straightforwardly expanded to 3D, working with one single encoding matrix. However, the latter will
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usually be too large for direct inversion and require iterative treatment. Alternatively, and more efficiently, algebraic processing can be restricted to 1D sub-problems, reconstructing full k-space lines through the origin from pairs of acquisitions with opposite gradient direction. The resulting full
3D image [7].
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set of radial k-space data is then merged by standard 3D gridding (see Section 4.5) to form the final
Data extrapolated to the gap in the described manner reflect the actual T2 decay during the dead time.
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In particular, the value at k = 0 is not affected by transverse relaxation. Hence, although these data points are not actually acquired, they correspond to signal obtained as if sampling was started at TE =
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0, which has led to the “ZTE” notation. 5.4
Bandwidth limitations
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There are limitations to the gap size and hence the signal bandwidth that can be handled with the algebraic approach. These arise from deteriorating conditioning of the encoding matrix as gap size increases [94]. Ill-conditioning boosts error in the signal and the signal model, causing three kinds of image distortion: 1) enhancement and spatial correlation of noise, 2) aliased and amplified out-ofband signal, and 3) intra-voxel oscillation of the spatial response [95]. Usually, the first two are dominant in practice and will therefore be described in more detail below. As the reason for illconditioning is data missing in central k-space, all image distortions described are of low-frequency nature and may not be distinguished easily.
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1) Noise in common Cartesian k-space raw data exhibits constant variance which leads to the usual uncorrelated appearance of noise in MR images. In contrast, in reconstructed ZTE data, noise variance increases strongly towards the k-space center, thus causing noise correlation in image space. Therefore, at large gaps, noise in ZTE images can have an unusual artifact-like appearance (Fig. 12). Its random nature is only observed by varying shapes in the image resulting from different noise input
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data. Importantly, ZTE image reconstruction benefits from intrinsic averaging in central k-space
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associated with the sampling density of 3D radial data increasing strongly at small radii. Therefore, distortions appearing in 1D projections do not necessarily translate into deterioration of 3D image
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quality.
2) Out-of-band signal violates the fundamental assumption of ZTE reconstruction in that it stems from
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locations outside the FOV. As a consequence, such signal will be assigned into the FOV, corresponding to aliasing. Furthermore, related to the bad conditioning it will appear strongly amplified in the image. This is a serious practical issue as ZTE acquisition not only captures the
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desired short-T2 signal of the object of interest but also unwanted contributions from nearby hardware parts, in particular RF coils (Fig. 13). The straightforward solution of targeting a larger FOV
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including all detected signal is of limited scope as it results in larger bandwidth, i.e., larger gap size, as well as longer scan times and larger data sets. Better ways of reducing such background signal are suppression techniques [7, 98], image subtraction [99], and designing RF coils with negligible content
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of the observed nucleus (see Section 9.3).
Within these boundaries and depending on setup and sequence parameters, practical gap sizes in ZTE experiments are usually limited to approximately 2 - 3.5 DW, where DW = 1/BW is the Nyquist dwell
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time for the bandwidth (BW) in the FOV. Using ZTE-optimized RF hardware, this can nevertheless translate into BW as high as 500 kHz [93], which is the key to high-resolution short-T2 MRI (see
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Section 3.1). 5.5
RF excitation
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RF excitation in ZTE imaging plays a particular role as it occurs while the radial encoding gradient is already on (Fig. 6d). To ensure consistency between data of different radial directions, the spectrum of the excitation pulse must be largely uniform within the bandwidth spanned by the gradient across the object. In the basic ZTE sequence, a short block-shaped, so-called hard pulse is used. For greatest uniformity the pulse duration should not exceed approximately DW/3 [100]. However, in practice durations up to about DW can also provide useful results. For bandwidths of hundreds of kilohertz this limits hard pulse durations to a few microseconds, which is much shorter than RF pulses commonly applied on
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clinical MR scanners. With the maximum B1 amplitudes usually available on such systems, flip angles of only a few degrees can be obtained (see Section 9.3). On the other hand, as little time is spent on RF pulses and gradient switching, the ZTE sequence can be operated with very short TR (see Section 7.1), even below 1 ms [93], thus enabling maximum steady-state magnetization even at very small flip angles.
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If flip angles should nevertheless be increased due to low B1 or very high bandwidth, or to create T1
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contrast, three options have been proposed, namely 1) longer hard pulses with intensity correction, 2) short modulated pulses, or 3) long modulated pulses with simultaneous excitation and acquisition.
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1) By increasing the hard pulse duration, its bandwidth will be reduced. For durations above
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DW, the main lobe of the Sinc-shaped pulse spectrum exhibits relevant variation in the FOV. Hence, larger flip angles are primarily obtained in the central FOV, which may still be useful if this coincides with the location of interest. Inconsistencies in signal intensity between
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projections can be reduced by correcting for the pulse profile [101]. This approach is limited by singularities above durations of 2DW and by variations of local steady-state magnetization
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over time.
2) The hard pulse is replaced by an RF excitation with both modulated amplitude and frequency [102, 103]. If designed appropriately, such pulses enable high spectral uniformity at high flip angle efficiency and thus outperform hard pulses of equal duration (Fig. 14). Unlike hard
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pulses, their length is not limited by uniformity considerations but by the maximum acceptable dead time (see Section 5.2). 3) Further extending modulated pulses may lead to intolerable dead times. To reconcile long
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pulses and short dead time it has been suggested in the SWIFT (sweep imaging with Fourier transformation) technique to overlap RF excitation and acquisition [104, 105]. This is indeed possible by alternating the two RF operation modes rapidly (Fig. 15), thus offering enhanced
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degrees of freedom in pulse definition. However, frequent switching comes at the expense of
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reduced acquisition duty cycle and thus SNR efficiency. Furthermore, feasible bandwidth is ultimately limited by the inverse of the dead time [79]. To image large bandwidths at reduced switching rate, the SWIFT approach has been expanded to a multi-band mode [106]. Alternatively, technically challenging approaches have been proposed where excitation and acquisition are performed truly simultaneously without gapping (see Section 9.3).
Apart from short hard pulses, all excitation options listed above generate some kind of spectral nonuniformity which needs to be considered during image reconstruction [101, 103, 104, 107]. A potential implementation obstacle may be the high temporal resolution required for creating highbandwidth pulses which is not always available on clinical scanners (see Section 9.4). A more
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fundamental limitation is posed by the large power deposited by high-bandwidth excitations, leading to potentially excessive specific absorption rates (SAR). 5.6
Comparison with UTE
In summary, the ZTE option exhibits the following advantages (+) and disadvantages (-) with respect
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to UTE imaging: Higher short-T2 sensitivity and resolution
+
Greater resilience against off-resonance effects
+
Greater robustness against eddy currents, no trajectory calibration required
+
Silent operation
−
High-bandwidth RF excitation: higher SAR and tighter limits on flip angles
−
Bandwidth limited by dead time
−
Only 3D
−
Limited intrinsic contrast
−
Higher sensitivity to background signal
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6 Intermediate techniques
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Various modifications of the basic sequences discussed so far in Sections 3.3, 4, and 5 have been proposed to mitigate some of their limitations. This results in the intermediate techniques indicated in
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the classification scheme of Fig. 5. They implement transitions on the two axes corresponding to concept B (vertical axis) concerning phase and frequency encoding and concept C (horizontal axis)
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concerning the succession of RF excitation and gradient switching. Intermediate techniques are obtained by either combining data obtained with different techniques or by merging opposed concepts within a sequence.
Filling the ZTE dead-time gap
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6.1
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Gap size limitations in algebraic ZTE have mainly been tackled by acquisition of additional data in the k-space center: 1) Sampling at k = 0 by additional acquisition in the absence of a gradient improves the conditioning [95, 108, 109]. However, this data point will have a finite TE equaling the dead time, thus leading to considerable signal decay for tissues and materials very short T2 values. This introduces an inconsistency with the bulk data which leads to image artifacts for badly conditioned algebraic reconstruction. 2) The previous approach is expanded in the PETRA (pointwise encoding time reduction with radial acquisition) technique [110] by filling the entire spherical void in k-space with
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Cartesian SPI data (Fig. 16a). In this way, arbitrary gaps can be handled, though at the expense of scan time and SNR efficiency. The effect on these two parameters increases steeply with gap size. 3) Alternatively, in the technique known as “water- and fat-suppressed proton projection imaging” (WASPI) [111], the dead-time gap is filled with radial data acquired at a gradient
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strength reduced by a factor 2 × dead time / DW (Fig. 16b), thus achieving Nyquist sampling
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in the very center also. By adapting the angular density to the radius given by the gap size, the additional scan time is usually less than for PETRA.
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The latter two options change the weighting obtained in k-space associated with T2 decay. The related effect of on image quality can be assessed via the PSF (Fig. 17). The PSFs show that compared to
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ZTE, the main lobe width is reduced for both options but the PETRA technique offers relatively benign side-lobe behavior whereas WASPI is prone to increased side lobes. It has been found that this tendency can lead to artifacts in the presence of components with very short T2 on the order of the gap
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size, stemming, e.g., from hardware parts (see Section 9.3) or solid tissues [109].
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4) In the hybrid filling (HYFI) technique [112], radial data from a WASPI acquisition with multiple gradient strengths and a Cartesian SPI acquisition in the very center similar to PETRA are combined. This approach enables improved PSF quality as compared to WASPI with less scan time increase than required for PETRA.
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In the techniques described above, data missing in k-space is complemented by performing additional gradient encoding. In a different approach following the principles of parallel imaging [97],
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supplementary information is obtained by detecting the MR signal with a receiver coil array. During image reconstruction, independent data from multiple coils facilitates deriving missing from existing neighboring data, thus improving the conditioning associated with large gap size [113]. A challenge of this approach is the need to solve the full 3D inverse problem for image reconstruction, differently
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from the simplified procedure outlined in Section 5.3. Finally, the encoding potential of RF pulses
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may be exploited to reduce the effective gap size [114]. 6.2
Reducing the excitation bandwidth
Another class of intermediate techniques departs from single gradient ramps either before or after the RF pulse by performing excitation under a lower gradient which is then ramped to full strength for rapid encoding. Such techniques have been proposed as gradient-modulated PETRA [115] or RHE (ramped hybrid encoding) [116] as well as gradient-modulated SWIFT [117]. As an advantage, RF excitation can be performed at lower bandwidth, thus relaxing flip-angle and SAR constraints. On the
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other hand, gradient ramping per se and its fidelity affect image quality in analogy to the UTE case (see Sections 4.1 and 4.2). 7 Further sequence aspects 7.1
Repetition time
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The sequences that have been described above are usually operated with short TR and a steady state of
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the magnetization. In this way, the many repetitions required for 3D acquisition can be performed within useful scan times. Assuming complete spoiling of transverse magnetization after each TR,
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maximal transverse steady-state magnetization ' and thus optimal SNR are obtained by excitation
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at the Ernst angle ( [118]. For ) * , which holds for most short-T2 imaging scenarios, both ( and ' are approximately proportional to ) (see Section 12.2). ' + ) implies that, when operating in the Ernst-angle regime, changing TR has no effect on SNR efficiency [119]. For example,
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reducing TR by a factor of two while doubling the number of averages will yield the same SNR at equal scan time. Furthermore, ( + ) implies that changing TR does not affect the SAR generated
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by the excitation.
To achieve short TR, gradient spoiling must be implemented in an efficient way (see Section 0). Particularly fast repetition is possible with minimal gradient switching as permitted by SPI (see Section 3.3) and ZTE (see Section 5.1). For these techniques, as an additional benefit, reducing TR
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helps approach the Ernst angle which may otherwise not be reached (see Section 5.5). Hence, in highbandwidth ZTE scanning, TR may range down to only hundreds of microseconds whereas UTE
7.2
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techniques are often operated with TR values of several or even tens of milliseconds. Gradient spoiling
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To obtain a strict steady state, the same net gradient area must be applied in all dimensions for each TR [120]. Hence, in radial sequences each readout gradient must be rewound and gradient spoilers
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must be applied along dimensions fixed in physical space to destroy unwanted coherences [38]. Therefore, despite sub-millisecond encoding times, TR often amounts to several milliseconds. To reduce TR, pseudo-steady-state operation can be employed by leaving out spoilers and instead randomizing gradient areas [65] and/or projection orders [7, 108]. The latter is common particularly in ZTE and SPI sequences where switching gradients is usually avoided to minimize TR (see Section 7.1) and to keep their near-silent nature. Spoiling is then performed by extending the readout gradient along with suitable projection ordering.
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7.3
Reducing scan time
3D radial scanning is relatively time-consuming, particularly at large matrix sizes. To fulfil the Nyquist criterion for center-out encoding, a factor of π more encoding steps are required compared with a conventional Cartesian sequence. Ways of reducing the scan time include generalized parallel imaging schemes [97, 121], angular undersampling [122] (see Section 7.4), and compressed sensing
F
[123]. Furthermore, the k-space area covered per TR may be increased by using other center-out
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encoding schemes such as, e.g., spiral- [124], or cone-shaped [125-128] trajectories, within the limits given by gradient specifications and targeted T2. Another modification proposed for the UTE gradient
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shape aims at SNR optimization by adapting the k-space speed to compensate for the non-uniform
7.4
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density of radial encoding [129]. Geometry
The specific way of performing spatial encoding with the described techniques has implications for
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scan geometry.
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1) For both pure phase and radial frequency encoding there is no particular readout dimension in which bandwidth-limited filtering can be applied to restrict the FOV. Although possible for individual 1D projections in the radial techniques, such filtering would lead to inconsistent 3D data similar to selective RF excitation (see Section 5.5). Hence, in all (usually three)
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spatial dimensions a FOV given by object shape and the sensitivity range of the RF coils needs to be fully encoded to avoid aliasing. Thus potentially more encoding effort is required than with conventional sequences.
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2) As opposed to the discrete nature of the PSF in Cartesian imaging, the PSF for radial encoding exhibits a dense pattern of low-amplitude streaks. Therefore, undersampling causes streak artifacts of intensities much lower than with Cartesian sampling, and reducing the
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angular distance of the radial spokes by factors 2-4 leads to relatively benign and often
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tolerable artifacts.
3) For isotropic encoding, the FOV in 3D radial techniques is spherical. A simple adaptation to the object shape can be achieved by individual scaling of the gradients per dimension, resulting in an ellipsoidal FOV and anisotropic resolution. Disentangling these two parameters is possible with a generalized approach to radial encoding [130]. 4) Creating FOV off-centers with radial encoding is also more involved than with Cartesian techniques. For off-center excitation under a gradient, as in the SPI and ZTE sequences, the RF frequency must be adjusted to the respective gradient direction [7]. Similarly the receiver frequency needs to be set individually. In UTE encoding, due to gradient ramping, the receiver demodulation frequency must be changed also during the readout, which may be a
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limitation with some spectrometers. Alternatively, after sufficiently broadband acquisition, off-centers are applied by data demodulation during image reconstruction. 7.5
Contrast
The utility of medical images hinges on contrast between tissues types or conditions. MR image
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contrast is controlled by changing sequence parameters (e.g., excitation flip angle) or adding magnetization preparation (e.g., inversion), and further modified by combining multiple images (e.g.,
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by subtraction). In principle, in short-T2 imaging, the same concepts can be applied to generate
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contrast. However, a specific requirement is the separation of short-T2 signals from the often contrast generation in short-T2 techniques:
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dominant signals of long-T2 tissues. Furthermore, some specific considerations apply concerning
1) With negligible echo times, all tissues contribute according to their spin density and steadystate magnetization, often leading to images with relatively low intrinsic contrast.
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2) The effect of RF pulses on magnetization depends on their duration relative to T2 [52, 131] which needs to be taken into account for excitation, refocusing, and preparation. This may be
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limiting in some cases but can also be utilized to achieve T2 selectivity [76, 131-135]. 3) T2* weighting can be introduced in the UTE and SPI sequences by increasing TE, an option that obviously does not exist for the ZTE approach. A related effect may be achieved by
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reducing the encoding duration [7]. However, this approach emphasizes T2* decay at higher spatial frequencies.
4) T1 weighting using large flip angles is most feasible for the UTE technique but with long
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pulses, T2 decay during the pulse may need to be considered (see Section 4.3). For ZTE and SPI excitation, restrictions to flip angles exist associated with RF amplitude, dead time, and SAR (see Section 5.5).
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5) Most modes of magnetization preparation require more time than the TR values typical of the steady-state sequences described above. Hence, efficient performance requires a segmented
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implementation as employed in the Cartesian technique called “magnetization prepared rapid gradient echo imaging” (MP-RAGE) [136]. Emphasizing contrast by specific acquisition ordering as in the Cartesian case is not possible for basic UTE or ZTE scanning, as in all TR intervals equal parts of k-space are acquired. However, SPI and some of the hybrid techniques (see Section 0) enable such contrast manipulation. For the potentially SAR-intensive ZTE and SPI sequences, adding preparation pulses may be limited.
6) Subtraction images based on multiple gradient echo acquisitions with different TE can provide excellent T2* contrast. However, on radial trajectories, refocusing is sensitive to eddycurrent-induced artifacts due to concentration of errors in central k-space and requires
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correction measures [64, 66, 137-139]. For multi-echo ZTE scanning, eddy currents can be mitigated by employing the echo-shifting principle [140] to reduce gradient switching [141, 142]. 7) With zero or ultra-short TE, the image phase will be less affected by off-resonance than in conventional echo-based sequences. However, phase accrued during RF excitation and
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gradient encoding still contains information about local susceptibility or chemical shifts that
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can be used as a source of image phase contrast [143-145].
Within these constraints, contrast can be generated in short-T2 imaging. Several approaches have been
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proposed to separate short- from long-T2 signal, in particular to generate images with short-T2 signal only [38, 80, 132, 133, 146-158]. This results in so-called positive contrast for short-T2 tissues as
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opposed to classical T2-weighted images where these tissues are identified by the absence of signal (Fig. 18). Moreover, contrast within short-T2 components has been obtained by weighting by or quantification of T2 and T2* [53, 81, 159-163], T1 [162, 164-169], T1ρ [170, 171], chemical shift [150,
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172, 173], or magnetization transfer [167, 174-178]. Importantly, the last option requires suitable measures to separate true magnetization transfer from direct saturation which is stronger for lines
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broadened by rapid transverse relaxation [175]. 8 Artifacts and post-processing
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In this section, further aspects of image quality are discussed in addition to the resolution limitations and sequence-specific artifacts described in Sections 3-5. Off-resonance
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8.1
In radial images, the off-resonance effects due to chemical shifts, global field variations, or local susceptibility gradients deviate from the related signal shifts observed in Cartesian gradient echo
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techniques. With the center-out acquisition scheme often used for short-T2 MRI, off-resonance phase is accrued to form a cone-like shape in two-dimensional k-space or a corresponding structure for the
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3D case. This leads to blurring or ringing depending on the resonance offset. The transition between the two effects occurs for offsets between 1 and 2 PWB, where the pixel bandwidth of an image is defined as PBW = BW / N and N is the number of pixels in the FOV. To reduce off-resonance due to B0 non-uniformity, shimming is usually applied. The generic means of
mitigating off-resonance artifacts of any origin is using strong gradients, i.e., large PBW, as employed anyhow for encoding short-T2 tissues. Complementary, off-resonance correction can be applied to the data based on B0 mapping, using conjugate phase reconstruction [179-182] or iterative approaches [183, 184]. As in other sequences, a range of off-resonances values within a voxel leads to signal loss
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until the start of data acquisition and signal decay during the readout, which can only be mitigated by shortening TE and increasing encoding gradients, respectively. A particular effect in vivo is created by the chemical shift between water and fat of approximately 3.5 ppm. This corresponds to hundreds of Hertz at usual field strengths which is on the order of commonly used PBWs. If both water and fat contribute to the MR signal, the demodulation frequency
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chosen for data acquisition and/or image reconstruction can be on resonance for only one of them,
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leading to image blurring for the respective other component. This has a strong impact on the appearance of tissue interfaces and thus image contrast as demonstrated in Fig. 19. It is important to
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keep these effects in mind in practice because artifacts such as those seen in the figure might be misinterpreted. Besides using higher PBW, such effects may be avoided by suppressing either fat or
8.2
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water using magnetization preparation or separation techniques [144, 185-190]. Resolution enhancement
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As outlined in Section 3.1, resolution in MRI is limited by apodization in k-space caused by T2 decay affecting the acquisition range. However, in analogy to time-domain corrections in spectroscopy
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[191], such apodization can be countered by suitable correction factors in k-space, corresponding to a filter operation in the image domain. In particular, T2 filtering can be applied which compensates for exponential T2 decay (Fig. 20):
',!- &
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,!- &
(5)
When setting ',!- & of a given target tissue, the spatial resolution of this tissue will not be
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limited by the encoding duration. However, this tempting approach has two serious limitations: 1) While resolution will be improved for all . ',!- &, high spatial frequencies will be overemphasized for any / ',!- &, leading to ringing artifacts. For samples with excessive
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T2 range, the spatial variation of T2 can be included in the signal model [192]. However, this
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approach requires additional T2 mapping and more extensive computation.
2) Multiplication by the exponential in Eq. (5) boosts not only high-k signal but also noise, at the expense of image SNR.
Therefore, although moderate T2 filtering may be a useful means of optimizing image appearance, it cannot replace sufficiently rapid encoding in short-T2 MRI. 9 Hardware As discussed in preceding sections, imaging of tissues with short T2 differs from common MRI in terms of acquisition strategies and sequence timing. As a consequence, it alters requirements on the
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instrumentation used. In several respects, short-T2 imaging calls for enhanced hardware performance while in others it demands lower or different specifications. For some short-T2 techniques, dedicated hardware is not essential. Nevertheless, for best results and for conquering ever shorter T2, tailored hardware is clearly beneficial if not indispensable. Magnet
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9.1
Today, the strength of the static B0 field of most clinical MRI systems is 1.5 T or 3 T, whereas in
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numerous research systems it is 7 T and reaches up to 10.5 T [193]. B0 affects mainly three aspects:
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1) Baseline sensitivity: As usual in human MRI setups, the SNR increases approximately linearly with B0 [194]. SNR is particularly critical for tissue components with short T2 due to
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often low spin density and long T1 (see Section 2.5) as well as short, high-bandwidth acquisitions. Thus, with regard to baseline sensitivity, high B0 is favored. 2) Relaxation times: For field strengths and tissues of interest, T1 increases along with B0 while
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both T2 and T2* decrease [13], eating up part of the sensitivity gain. In turn, at high field, shorter T2 calls for potent short-T2 techniques and longer T1 entails lower optimal flip angles.
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3) B1 efficiency: The power efficiency of generating B1 declines as B0 increases, approximately according to 0 120 0 [195, 196] with 120 being the power absorbed in the exposed tissue. This dependency renders RF power demands and SAR at high field challenging
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despite slightly reduced flip angle requirements.
Optimal choice of the field strength accounts for all of these aspects and also depends on the tissue
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and application of interest.
Concerning B0 uniformity, short-T2 techniques pose reduced demands since they use strong gradients and high-bandwidth acquisition. Therefore, per-subject shimming is often not required. It still tends to
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be relevant, however, when weaker or ramped gradients are employed during data acquisition, such as in WASPI, UTE imaging, or gradient-modulated variants (see Section 6.2). Furthermore,
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magnetization preparation may require more uniform B0 than data acquisition (see Section 7.5). 9.2
Gradient system
The gradient system [197] has two main types of components, amplifiers and coils, which together produce dynamic field gradients according to input waveforms. 1) Gradient strength: The maximum gradient strength G is a key parameter as it determines kspace speed, which competes with signal decay. For purely frequency-encoded techniques (e.g. UTE or ZTE imaging) the latter causes approximately exponential apodization (Figs. 8 and 17). To actually achieve a given nominal resolution %&, the readout gradient must
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approximately obey 3 $ %& [37, 38]. For example, for T2 = 500 μs an actual resolution of 1 mm can be obtained with G = 23.5 mT/m, which is feasible on most clinical systems, providing G is in the range 20 - 80 mT/m. However, T2 = 50 μs requires the extreme value of G = 235 mT/m or will yield only 10 mm resolution at the lower G. For purely phaseencoded techniques (e.g. SPI), resolution is not limited by T2, but large G is still important to
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limit TE for best SNR (see Section 3.3 and Fig. 17).
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2) Slew rate: Techniques that set readout gradients before RF excitation (e.g., SPI and ZTE) may operate gradients quasi-continuously with minimized switching (see Section 5) and thus do
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not depend on high slew rates. Still, faster gradient switching permits shorter TR, which is associated with smaller optimal flip angle, higher acquisition duty cycle, and less overall scan
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time (see Section 7.1). In contrast, when RF excitation is performed first (e.g. in UTE), the speed of gradient ramping is a key determinant of resolution and robustness against offresonance (see Section 4.1, Fig. 8). Slew rates in commercial scanners range up to 200
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mT/m/ms in combination with a gradient strength of, e.g., 40 mT/m. Importantly, to efficiently use an increase in gradient strength, the slew rate must grow quadratically [119].
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3) Duty cycle: Quasi-continuous operation of ZTE- and SPI-like sequences means that gradients are always on, often at constant net magnitude, but with changing direction. This can only be accomplished if full duty cycle is provided at the applied high strength, which is a rather
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uncommon design criterion for gradient systems. Although the same techniques can also be run non-continuously with gradient switching and reduced duty cycle, the continuous mode is clearly preferable as it permits shorter TR and strongly reduces eddy currents and acoustic
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noise. For methods with intrinsic gradient switching (e.g. UTE imaging), high duty cycle is less essential but still helps minimize TR.
4) Amplifier: Requirements for the gradient amplifier arise from the three specifications
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discussed above. Gradient strength and slew rate are directly related to maximum current and voltage, respectively. Some amplifiers offer added flexibility by permitting multiple
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combinations of peak current and peak voltage [198]. Operation at full duty cycle requires that the current necessary for a desired gradient strength can be provided continuously along with the corresponding power for given ohmic losses in the gradient coil.
5) Fidelity: The spatiotemporal fidelity of gradient fields is generally hampered by delays, finite bandwidth, eddy currents, and amplifier imperfections. This affects image quality primarily in UTE techniques (see Section 4.2) and when additional echoes are generated for creating contrast (see Section 7.5). To a certain degree, these effects can be reduced by delay calibration and pre-emphasis. If necessary, residual field errors can be measured and accounted for at the level of image reconstruction (see Section 9.5).
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6) Peripheral nerve stimulation: In vivo, gradient switching must avoid peripheral nerve stimulation (PNS) induced by rapid, large-amplitude field changes [199]. In this respect, short, localized gradient coils are advantageous over whole-body gradients [197]. Short-T2 techniques with minimized gradient switching are not a problem with regard to PNS. Sequences with switching need more attention but are typically less prone to PNS than, e.g.,
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echo-planar imaging with long trains of shortly spaced rapid field changes.
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In summary, fulfilling all of the listed criteria at the same time is not trivial but it is possible with
9.3
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dedicated designs, thus also enabling extreme applications [200]. RF chains
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Short-T2 MRI relies on high-bandwidth, sub-millisecond acquisitions and thus also alters requirements on RF hardware. For UTE imaging, the factors to consider are rapid T/R switching and avoiding background signal. For ZTE and SPI techniques, which require excitation during readout
power is needed (see Section 5.5).
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gradients, these demands are further increased, and in addition, efficient RF transmission at high
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1) Amplifier: For ZTE and SPI, the bandwidth of RF amplifiers must be sufficiently large to enable high-bandwidth excitation using either short hard pulses of duration in the μs range or frequency-swept pulses with large sweep width (see Section 5.5). Moreover, rapid amplitude
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changes at the beginning and end of these pulses call for sub-microsecond rise times. In analogy to gradient amplifiers, attenuation in transition bands may be overcome by preemphasis as far as permitted by power margins. Moreover, accurate realization of modulated
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pulses requires good linearity which may also be improved by means of calibrated compensation. High-power transmission is particularly important to achieve adequate flip angles by hard-pulse excitation. As an example, when using an RF coil providing a typical B1
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efficiency of 20 45 67, obtaining a flip angle of 5° with a hard pulse of 2 μs duration
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requires 2.9 kW. Furthermore, the available duty cycle must permit application of the desired pulses (few to hundreds of microseconds) at short TR (potentially sub-millisecond, see Section 7.1) and for long scan times (potentially tens of minutes). Amplifier ring-down after the RF pulse causes artifacts in high-bandwidth images when time constants are similar to T/R switching times [201]. To suppress this ring-down sufficiently, blanking after transmission must be fast and achieve a high degree of isolation in addition to the attenuation provided by T/R switching or de-tuning (see below).
2) Coils: MRI generally relies on either a single coil for both RF transmission and detection or separate devices for the two purposes. The latter setup usually consists of a receiver array
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near the subject to maximize SNR and a surrounding transmit coil. However, those short-T2 techniques with high B1 requirements call for particularly high transmit efficiency, which can favor T/R coils, in particular T/R arrays. A further advantage of such single-coil setups are reduced space requirements along the radial direction. On the other hand, for safety reasons, the distance to the subject needs to be larger than for a receive-only coil. Furthermore, pre-
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amplifier decoupling cannot be implemented. On the receive side, as usual, coil arrays offer
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SNR benefit [202] and parallel imaging capability [97]. In ZTE techniques, complementary spatial encoding by array detection also holds potential to support filling the k-space gap (see
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Section 6.1). Similar to RF amplifiers, the bandwidth of RF coils must be large enough to generate and detect, at sufficient uniformity, suitable excitation pulses and resonance signals,
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respectively. If not provided by conventional coils, dedicated designs are required. For residual imperfections, appropriate pulse compensation or data correction procedures may be employed. A further design criterion arises from the original aim to capture short-T2 signals
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from the subject. In addition to the latter, short-T2 detection is prone to capture MR signals also from hardware parts, which conventional MRI does not pick up due to their rapid decay.
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This issue is most prominent for materials close to coil conductors where B1 is high (see Section 5.4). Ideally, such contributions are avoided by choosing materials that do not contain the targeted nucleus. For 1H imaging this is achieved with, e.g., glass or fluorinated polymers such as polytetrafluoroethylene (PTFE) [82, 203-205].
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3) T/R switching: In pulsed MR, the receive chain must be sufficiently isolated from the transmit path in both operation modes. During transmission, destruction and preferably even saturation
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of receive pre-amplifiers must be prevented. During reception, residual ring-down of the power amplifier must be blocked. Depending on the coil configuration, either switching between the two operation modes by means of a T/R switch or de-tuning of the transmit coil and tuning of the receive coil are required. The particular challenge for short-T2 techniques is
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to perform this rapidly, i.e., in the low microsecond range as opposed to several tens of
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microseconds in commercial systems, to minimize the RF dead time [206, 207] (see Section 5.2). To avoid image artifacts, switching transients must have decayed to a level well below that of the MR signal when data is acquired. Rapid switching is particularly challenging at high peak and average power. It is typically based on positive-intrinsic-negative (PIN) diodes [208], but has also been achieved with transistors [208, 209] or micro-electro-mechanical systems (MEMS) [210]. In a recent design based on PIN diodes, switching transients are cancelled by symmetric biasing [211].
4) Simultaneous operation: A tempting alternative to RF switching is simultaneous transmission and reception [212-214]. With this approach, dead time overheads are avoided and sequences
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are simplified. Simultaneous transmission and reception was common practice before the advent of pulsed NMR, albeit in rigid setups and at much lower frequencies. For in-vivo imaging at present-day MR frequencies, sufficient isolation between transmit and receive paths is difficult to achieve without T/R switching or detuning. Isolation between the two RF paths can be achieved by geometric coil decoupling, additional decoupling by passive
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electronics, and active cancelation of transmit leakage into the receive path. However, the
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robustness of such setups is limited by residual variation in RF conditions, particularly by patient motion, which causes subtle changes in RF loading. Feedback control of cancelation
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has been proposed to address this limitation [215, 216]. In an alternative approach, isolation is facilitated by sideband modulation, which permits performing transmission and reception in
9.4
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different frequency bands [217]. Console
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The console of an MRI system includes a sequencer generating RF and gradient waveforms, a receiver for data acquisition, and a reconstructor that calculates images from raw data. This
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architecture is the same for short-T2 imaging. However, for high performance in the short-T2 mode, the components should meet more stringent specifications. 1) Sequencer: The temporal resolution of the sequence description should be 1 μs or below. The
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definition and generation of RF pulses must be more flexible than in common clinical MRI systems, including a wider range of pulse durations (1 μs to tens of milliseconds) and TR (down to 100 μs, see Section 7.1). To perform off-center imaging with ZTE or SPI (see
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Section 7.4), per-shot adjustment of the excitation carrier frequency is necessary. The same is true for the receiver with any radial technique.
2) Receiver: The feasible acquisition bandwidth must cover the signal bandwidth of up to the
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megahertz range. Even higher acquisition bandwidth is needed for oversampling to enable finite-support extrapolation (see Section 5.3). In conjunction with high spatial resolution and
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array detection, 3D acquisition with oversampling yields very large amounts of data that need to be handled and stored. High bandwidth and short TR render these data challenging also in terms of throughput. Short-T2 imaging benefits particularly from flexibility at the level of digital filtering, where short, customized, and multi-rate filters [218] permit minimizing RF dead times without bloating the raw data (see Section 5.2).
3) Reconstructor: 3D radial data sets tend to be very large and do not readily lend themselves to splitting into smaller portions for separate processing. Therefore, a computer used for image reconstruction should be equipped with ample memory to avoid time-consuming swapping on the hard disk. For example, the raw data for image matrix size 256 acquired with full Nyquist
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angular sampling, 4-fold radial oversampling, and 8 channels amount to approximately 13 GB. For 3D gridding reconstruction, about 8 times that amount of memory needs to be allocated. Reconstruction time is longer than in standard Cartesian imaging and is determined by the numerical complexity of algebraic parts, density correction, gridding interpolation, and FT, and by the number of cycles in the case of iterative reconstruction (see Sections 4.5 and
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5.3). For speed-up, reconstruction may be partly parallelized and distributed to different nodes
Calibration tools
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9.5
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on a single computer or a cluster of computers, or on graphical processing units (GPUs).
All magnetic fields – static, audio- and radio-frequency – experienced by the spins contribute to
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encoding of the MR signal. Hence, for proper image reconstruction, the spatial and temporal behavior of these fields must be accurately included in the underlying signal model. However, encoding fields are produced with finite accuracy by the respective hardware, i.e., the magnet, gradient subsystems,
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and RF chains. To determine actual fields and field evolutions, either the MR system itself or additional dedicated measurement hardware may be used to record field information at runtime or in a
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calibration step. Depending on the severity of deviations, the results can be used to improve image reconstruction or to pre-distort input waveforms for higher accuracy. 1) Gradients: Apart from the methods discussed in Section 4.2, spatial and temporal behavior of
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gradient fields per se as well as eddy current effects can be measured by a set of MR field probes [71]. A field probe is a small RF coil wrapped around an NMR sample and determines the variation over time of field strength at one position. Several such probes enable spatial
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expansion of the field dynamics in the imaging volume. Such information has been used to improve reconstruction of, e.g., UTE images [219, 220] and dual-echo ZTE data [139]. Instead of measuring field behavior for each protocol, by assuming linearity and time
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invariance, the gradient system may also be characterized by its impulse response [221], enabling field calculations for arbitrary gradient shapes.
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2) RF: Effective RF waveforms are subject to variable degrees of distortion caused by RF amplifiers, lines, cables, and coils. These are most easily observed with a pick-up coil placed near the sample and connected to an MR spectrometer or a spectrum analyzer. In this way, effective RF pulses as experienced by the spins are observed, including all disturbing effects in the transmit chain. Alternatively, field probes may be used in an advanced setup to monitor RF and gradient fields simultaneously [222], including information on spatial variation. Knowledge of effective RF waveforms is particularly important for the use of amplitude- or phase-modulated pulses (see Section 5.5).
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9.6
Summary
The hardware considerations compiled here give an indication of how dedicated systems for short-T2 MRI might differ from present-day instrumentation. For optimal UTE imaging, requirements are actually rather similar to those of conventional MRI. High B0 should be combined with high gradient performance in terms of amplitude, slew rate, and
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fidelity, and careful consideration of PNS. With regard to RF, in a similar manner to regular MRI,
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dense receiver arrays are desirable for optimal SNR. As one feature specific to the short-T2 scenario, a UTE system should offer rapid T/R switching to minimize TE. Integrated field monitoring equipment
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would facilitate accurate gradient characterization.
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In contrast, a system designed for ZTE imaging would differ more fundamentally. Due to the more limiting role of RF power considerations, the optimal field strength will be lower for ZTE imaging than for UTE. B0 field homogeneity is less critical in the ZTE scenario and may be traded off against
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magnet length and cost. Gradients should offer high amplitude and, critically, unlimited duty cycle. However, neither high slew rate nor gradient fidelity are particularly important, which will alter and
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potentially simplify coil design and amplifier considerations. Likewise, PNS is of little concern. On the RF side, the ZTE mode is more demanding, calling not only for dense array detection but also for high-power transmit capability, efficient transmission, very rapid T/R switching, and 1H-free coil front-ends. A ZTE console should offer superior temporal definition, whereas integrated RF
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monitoring would improve waveform correction and thus expand the range of feasible pulse modulation schemes.
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10 Applications
Most applications of short-T2 MRI relate to its ability to depict solid tissues. This is in contrast to conventional MRI, which depicts such tissues mostly by the absence of signal, relying on so-called
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negative contrast. By using dedicated short-T2 techniques instead, signals from solid tissues can be
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observed directly and thus with positive contrast. This approach has two major advantages: 1) Short-T2 tissues can be distinguished from actually void spaces.
2) Signal variation related to spin density or other sources of contrast (see Section 7.5) permits differentiation between different short-T2 tissues or between healthy and pathological tissues.
Table 2 lists several tissues and materials of interest along with T2 values as reported in the literature, indicating fields of application. With conventional MRI hardware, the short-T2 techniques described previously enable imaging of T2 values with several hundreds of microseconds at useful resolution. By employing dedicated RF and gradient hardware (see Section 9), T2 values below 100 μs also become realistic targets [201, 223].
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
1) Musculoskeletal (MSK) system: Short-T2 tissues are particularly prevalent in the MSK system, such as bone, tendons, ligaments, menisci and cartilage (Fig. 21a). The focus of many investigations is to image bone and quantify bone mineral density, either indirectly from bone water [49, 51, 53, 111, 151, 160, 168, 172, 176, 224-246] or directly from hydroxyapatite employing
31P-imaging
[6, 22, 244, 247-257], due to its potentially high diagnostic value,
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e.g., in osteoporosis. This application has also been expanded to bone substitutes [258].
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Furthermore, tendons [31, 38, 150, 170, 259-261], ligaments [49, 54], menisci [49, 51, 150, 170, 262], and cartilage [263-265] are of specific interest due to frequent injuries,
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inflammations, or wear.
2) Positron emission tomography (PET): The ability to depict bone can be utilized in devices for
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combined PET and MRI. PET requires attenuation correction for bone and other tissues which may be based on segmentation of short-T2 MR images [266-273]. The same approach can be applied for attenuation caused by hardware parts of the MRI scanner, in particular RF
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coils.
3) Pulmonary MRI with short-T2 techniques [47, 99, 108, 162, 167, 274-301] (Fig. 21b) has
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large potential, e.g., to improve specificity in early diagnosis of certain diseases (e.g., fibrosis) as compared to computer tomography (CT). For this application, a particular advantage of the radial short-T2 techniques is high robustness with respect to motion. 4) Iron oxide nanoparticles (IONP) are used as contrast agents for, e.g., cell labeling or tracking
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drug delivery. Typically, in MR images IONPs are observed as signal voids due to reduced T2*. Alternatively, short-T2 techniques enable imaging them with positive contrast, thus
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utilizing reduced T1 instead of T2* [7, 149, 165, 302-306]. In this way, their identification and localization can be improved.
5) Dental MRI using short-T2 techniques allows depiction of teeth with high quality (Fig. 21c). Application in vivo holds promise for improved early diagnosis of, e.g., caries lesions and
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benefits from dedicated hardware [307-315]. However, at present patient comfort, scan times,
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and cost cannot yet compete with standard X-ray techniques.
6) Myelin is a biomaterial that is essential in rapid electrical signal propagation along nerve fibers in white matter of the brain and the spinal cord. Therefore demyelination plays an important role in numerous neurological disorders. In a conventional approach, myelin is quantified indirectly via myelin water [316, 317]. Short-T2 techniques have the potential to directly visualize and quantify myelin but must rely on effective suppression of long-T2 tissue components [35, 147, 318-325].
7) Plaques in vessels due to calcification by arteriosclerosis reduce or obstruct blood flow in, e.g., peripheral, cardiac, or brain vasculature. Due to the short T2* of plaques, their assessment
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
is conventionally based on X-ray imaging. Short-T2 MRI offers a prospective alternative modality for plaque detection and specification [319, 326-329]. 8) Functional MRI (fMRI): MR techniques with ultra-short or zero TE are not intrinsically sensitive to blood-oxygenation-level dependent (BOLD) contrast. Nevertheless, they can be used for performing fMRI. Sensitivity to neuronal activation is either obtained by T2
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preparation [330] or based on other mechanisms, e.g., perfusion [331, 332]. As a particular
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advantage, the zero-TE techniques offer silent operation, avoiding brain activation related to acoustic noise generated by gradient switching.
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9) Electric property tomography (EPT): Spatial mapping of electric permittivity and conductivity by EPT has potential for various applications [333]. EPT is based on evaluation
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of the image phase. Techniques with negligible TE may contribute in this field due a reduced sensitivity of their image phase to off-resonance and eddy currents [334]. 10) Sodium MRI: Sodium is of interest due to its important role in many biochemical processes
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and is a potential biomarker in the diagnosis of diseases such as stroke, cancer, and osteoarthritis. The NMR-active isotope 23Na has spin 3/2, leading to bi-exponential relaxation
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behavior (see Section 2.3). T2 of the shorter component is typically best addressed with dedicated short-T2 techniques [126, 129, 335, 336]. 11) Materials: Apart from in-vivo uses, short-T2 MRI has large potential for materials research [337-340] and inspection as well as for applications in the food industry [341-343]. The
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ability to depict short-T2 materials is also instrumental in the non-destructive examination of mummies [344].
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12) Inhomogeneous B0: The use of minimized TE and high bandwidth, short-T2 techniques also provides considerable resilience against signal loss and displacement from strong macroscopic B0 (see Section 2.6). This kind of robustness improves imaging in the presence
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of tissue-air interfaces or metal implants [7, 345].
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11 Conclusions
In this article, we have reviewed the growing family of techniques for in-vivo MRI of tissues with rapid decay of transverse magnetization. Besides covering the variety of current approaches and implementations, this survey has taken a conceptual perspective, emphasizing fundamental challenges, derived principles, as well as commonalities and key differences among the various techniques. It has been shown that, despite their variety, short-T2 imaging techniques can be categorized according to a few fundamental design choices, leading to four basic sequence ideas, CTI, SPI, UTE, and ZTE. The Cartesian techniques CTI and SPI offer the attractive feature of enabling theoretically unlimited
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
spatial resolution even in the case of rapid signal decay. However, this option comes at the expense of poor SNR. Greater SNR can be achieved with the radial UTE and ZTE sequences, which are most widely used at this point. UTE imaging is feasible with conventional hardware but requires accurate gradient calibration. In comparison, ZTE approaches offer higher short-T2 sensitivity and resolution but are more demanding in terms of RF hardware and image reconstruction. The choice of technique
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thus depends on the targeted application and technical resources. Offering guidance for these
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considerations is one of the chief purposes of this article.
As an outlook, two roads can be foreseen in the emerging field of short-T2 MRI. First, clinical
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usefulness needs to be investigated for applications that are readily accessible with existing commercial scanners. These relate to tissues with T2 above approximately 300 μs, such as in MSK
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and pulmonary imaging. The basis for this task will be relatively established implementations of both UTE and ZTE variants, whereas further effort will be required to verify reliable contrast options. Second, the feasibility of imaging in vivo at T2 values below a few hundreds of microseconds,
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possibly down to tens of microseconds (as, e.g., in myelin) still needs to be explored and requires additional methodological and technical efforts. With regard to imaging methods, it is expected that
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there is still considerable potential in fusing and optimizing current approaches. Furthermore, on the technical side, advances in high-bandwidth capabilities of MR scanner hardware with respect to gradient strength and duty cycles, RF switching speed, console specifications, and proton-free
12 Appendices
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materials will pave the way for further discoveries in the world of short-T2 tissues.
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12.1 Implementation of the rho filter
For sequences with radial encoding, a suitable DCF is efficiently obtained by calculation of the rho filter. For center-out encoding schemes as used in UTE and ZTE imaging, it may be implemented as
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follows:
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89
:
89! ! !
(A1) :
! !
89! ! !
:
:
! !
(A2) :
(A3)
Here, ! ;;! counts the radial k-space positions starting from the center, k-values are given in arbitrary units, and d is the dimensionality of the data. Note that the DCF calculated according to Eqs. (A1-A3) is valid also for not being at as well as non-equidistant k-space positions.
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
12.2 Steady-state magnetization at short TR In a steady-state sequence with complete spoiling of transverse magnetization after each TR interval, the flip angle for maximized transverse steady-state magnetization ' is the Ernst angle [118] ( <=>>?@
(A4)
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with )
(A5)
O
O
leading to [346]
'
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For ) * the approximation [119] )
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#
( # &AABC
)
#
as well as
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leads to
)
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' #
)
(A6)
(A7)
(A8)
(A9)
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which by neglecting the -1 further simplifies to ' #
)
(A10)
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Hence at very short TR, both the Ernst angle and the signal are approximately proportional to ).
Funding
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
This research did not receive any specific grant from funding agencies in the public, commercial, or not-for-profit sectors. Declarations of interest MW None
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Declarations of interest KP
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Klaas Pruessmann holds a research agreement with and receives research support from Philips Healthcare. He is a shareholder of Gyrotools LLC and Skope
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Magnetic Resonance Technologies.
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[349] A.C. Seifert, S.L. Wehrli, F.W. Wehrli, BiDcomponent T2* analysis of bound and pore bone water fractions fails at high field strengths, NMR Biomed, 28 (2015) 861-872.
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[350] J. Du, E. Diaz, M. Carl, W. Bae, C.B. Chung, G.M. Bydder, Ultrashort echo time imaging with bicomponent analysis, Magn Reson Med, 67 (2012) 645-649. [351] R. Biswas, W. Bae, E. Diaz, K. Masuda, C.B. Chung, G.M. Bydder, J. Du, Ultrashort echo time (UTE) imaging with bi-component analysis: bound and free water evaluation of bovine cortical bone subject to sequential drying, Bone, 50 (2012) 749-755.
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[352] N. Kokabi, W. Bae, E. Diaz, C.B. Chung, G.M. Bydder, J. Du, Ultrashort TE MR imaging of bovine cortical bone: The effect of water loss on the T1 and T2* relaxation times, Magn Reson Med, 66 (2011) 476-482. [353] R.A. Horch, J.S. Nyman, D.F. Gochberg, R.D. Dortch, M.D. Does, Characterization of 1H NMR signal in human cortical bone for magnetic resonance imaging, Magn Reson Med, 64 (2010) 680-687. [354] R.A. Horch, D.F. Gochberg, J.S. Nyman, M.D. Does, Clinically compatible MRI strategies for discriminating bound and pore water in cortical bone, Magn Reson Med, 68 (2012) 1774-1784. [355] M. Robson, P. Gatehouse, T. He, D. Firmin, G. Bydder, S. Neubauer, Human in vivo imaging of phosphorus in cortical bone using ultrashort TE pulse sequences, In Proceedings of the 11th Annual Meeting of ISMRM, Toronto, Ontario, Canada, 2003, p.791 [356] L. Li, R.A. Kruger, P-31 NMR imaging of solid bone with solid echoes combined with refocused gradients, Phys Med Biol, 35 (1990) 1153-1158.
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[357] U. Grosse, F. Springer, T. Hein, G. Grozinger, C. Schabel, P. Martirosian, F. Schick, R. Syha, Influence of physical activity on T1 and T2* relaxation times of healthy Achilles tendons at 3T, J Magn Reson Imag, 41 (2015) 193-201. [358] V. Juras, S. Apprich, P. Szomolanyi, O. Bieri, X. Deligianni, S. Trattnig, Bi-exponential T2* analysis of healthy and diseased Achilles tendons: an in vivo preliminary magnetic resonance study and correlation with clinical score, Eur Radiol, 23 (2013) 2814-2822.
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[359] G.H. Filho, J. Du, B.C. Pak, S. Statum, R. Znamorowski, P. Haghighi, G. Bydder, C.B. Chung, Quantitative characterization of the Achilles tendon in cadaveric specimens: T1 and T2* measurements using ultrashort-TE MRI at 3 T, Am J Roentgenol, 192 (2009) W117-W124.
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[360] E.Y. Chang, W.C. Bae, S. Statum, J. Du, C.B. Chung, Effects of repetitive freeze–thawing cycles on T2 and T2* of the Achilles tendon, Eur J Radiol, 83 (2014) 349-353.
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[367] E. Ercan, P. Boernert, A. Webb, I. Ronen, Whole-brain tissue-based assessment of the ultrashort T2 component using 3D UTE MRI relaxometry, In Proceedings of the 20th Annual Meeting of ISMRM, Melbourne, Australia, 2013, p.4279 [368] Q. He, Y. Ma, S. Fan, H. Shao, V. Sheth, G.M. Bydder, J. Du, Direct magnitude and phase imaging of myelin using ultrashort echo time (UTE) pulse sequences: A feasibility study, Magn Reson Imag, 39 (2017) 194-199.
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[369] S.J. Fan, Y. Ma, E.Y. Chang, G.M. Bydder, J. Du, Inversion recovery ultrashort echo time imaging of ultrashort T2 tissue components in ovine brain at 3 T: a sequential D2O exchange study, NMR in biomedicine, 30 (2017) e3767. [370] S.J. Fan, Y. Ma, Y. Zhu, A. Searleman, N.M. Szeverenyi, G.M. Bydder, J. Du, Yet more evidence that myelin protons can be directly imaged with UTE sequences on a clinical 3T scanner: Bicomponent T2* analysis of native and deuterated ovine brain specimens, Magn Reson Med, 80 (2018) 538-547. [371] A. Ramani, A. Aliev, G. Barker, P. Tofts, Another approach to protons with constricted mobility in white matter: pilot studies using wideline and high-resolution NMR spectroscopy, Magn Reson Imag, 21 (2003) 1039-1043. [372] M.A. Schmidt, G.Z. Yang, P.D. Gatehouse, D.N. Firmin, FID-based lung MRI at 0.5 T: theoretical considerations and practical implications, Magn Reson Med, 39 (1998) 666-672. [373] M. Kveder, I. Zupančič, G. Lahajnar, R. Blinc, D. Šuput, D. Ailion, K. Ganesan, C. Goodrich, Water proton NMR relaxation mechanisms in lung tissue, Magn Reson Med, 7 (1988) 432-441.
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Glossary 1D: one-dimensional 3D: three-dimensional AC: autocorrelation
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AR: algebraic reconstruction
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B0: static magnetic field
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B1: radio-frequency magnetic field
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BLAST: back-projection low angle shot BOLD: blood-oxygenation-level dependent BW: bandwidth
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CT: computer tomography
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CTI: constant time imaging DCF: density correction function
EPT: electrical property tomography
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FID: free induction decay
fMRI: functional magnetic resonance imaging
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FOV: field-of-view
FT: Fourier transform
GM: gradient-modulated
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GPU: graphical processing unit
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HYFI: hybrid filling IONP: iron oxide nanoparticle MEMS: micro-electro-mechanical system MP mapping: multipoint mapping MP RAGE: magnetization prepared rapid gradient echo imaging MR: magnetic resonance MRI: magnetic resonance imaging
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MSK: musculoskeletal MTF: modulation transfer function NMR: nuclear magnetic resonance PBW: pixel bandwidth
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PET: positron emission tomography
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PETRA: pointwise encoding time reduction with radial acquisition
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PIN: positive-instrinsic-negative
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PNS: peripheral nerve stimulation PSF: point spread function PTFE: polytetrafluoroethylene
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RF: radio-frequency
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RHE: ramped hybrid encoding RUFIS: rotating ultra-fast imaging sequence SAR: specific absorption rate
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SPI: single-point imaging
SPRITE: single-point ramped imaging with T1 enhancement
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SNR: signal-to-noise ratio
SWIFT: sweep imaging with Fourier transformation T1: longitudinal relaxation time
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T2: transverse relaxation time
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T2*: apparent transverse relaxation time TE: echo time : acquisition range, or range of times across which data are collected
TR: repetition time T/R: transmit-receive UTE: ultra-short echo time WASPI: water- and fat-suppressed proton projection imaging
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
ZTE: zero echo time
Fig. 1: Dipolar coupling as an example of an orientation-dependent interaction. a) The dipolar field of one spin (red arrow) affects the total field at the location of another spin (blue arrow). The local field contribution %0E >?@ E depends on the orientation of the molecule as the spins keep their
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alignment with B0 upon molecular rotation, thus changing the position of the affected spin within the
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dipolar field. The relevant angle θ is the angle between the B0 field vector and the axis connecting the two spins (dashed line). Molecular tumbling changes the orientation of the molecule over time and the
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local field as seen by the spin fluctuates. b) Effects of dipolar coupling on NMR spectra. In a single crystal, all local lattices form the same angle with B0 (θ = 20° in this case), leading to a distinct peak.
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The second, symmetric peak arises from spin pairs in the same sample with anti-parallel instead of parallel mutual alignment (one of the spins in a) flipped). In a powder, all lattice orientations co-exist; thus the spectrum reflects all corresponding frequency shifts. The latter are weighted with @FG E
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according to their occurrence for an equal distribution of orientations, leading to the peculiar shape peaking at θ = 90°, the so-called Pake pattern. Again, both ways of alignment contribute. Note that
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the field shift vanishes at the so-called magic angle of θ ≈ 55°. In a hypothetical liquid with arbitrarily fast molecular rotation, the dipolar field is averaged to zero over time, leading to a single narrow line. In a real liquid, averaging also occurs, but in addition, the temporal fluctuations cause T1 and T2 relaxation, the latter leading to line broadening. Note that the spectra are scaled for presentation
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purposes. c) Free induction decay signals obtained for the spectra shown in b). The hypothetical liquid gives a constant on-resonant signal whereas the liquid with relaxation shows exponential decay. The
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off-resonant peaks of the single crystal lead to oscillations with constant envelope. The signal associated with the powder spectrum drops rapidly due to superposition of a wide range of frequencies. In this case, the observed loss of transverse magnetization is a non-random, reversible
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process and therefore not considered to be actual relaxation. Fig. 2: Relaxation in a nutshell. In the presence of an orientation- or position-dependent interaction
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(e.g. dipolar coupling, see Fig. 1), molecular tumbling leads to randomly fluctuating fields ΔB(t) at the spins’ locations. The temporal behavior of the random field is determined by the speed of the tumbling motion and characterized by the correlation time τc, whereas its amplitude depends on the strength of the interaction. Two typical random fields are shown with rapid (τc = 0.05 ns, red) and
slow fluctuations (τc = 0.4 ns, blue). The autocorrelation (AC) function calculated from the average of such a random field decays exponentially with the constant τc. Consequently, the Fourier transform (FT) of the AC function is a Lorentz function, also known as the spectral density J(ω) of the random field, i.e., the energy available at different frequencies, which will be required below. Relaxation of longitudinal magnetization involves changes of the populations of different spin states. This is
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
illustrated in the energy diagram drawn for an isolated pair of equivalent spins, where transitions between combined states occur with probabilities W0, W1, and W2 where the subscripts indicate the corresponding “step size” in the diagram. The transition probabilities scale with the spectral density taken at multiples of the Larmor frequency ω0. How they contribute to T1 is given by the particular interaction causing the relaxation. T1 relaxation is most efficient for τc ≈ 1/ω0. Transverse relaxation
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involves losses in coherences. Therefore, J(0) contributes and dominates T2 for large τc, leading to a
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continuous T2 reduction for slower motion. The dashed vertical red and blue lines indicate the relaxation times obtained for the example values of τc.
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Fig. 3: Principles of short-T2 MRI, illustrated by three different acquisitions performed on a phantom containing mineral oil with long T2 surrounded by pieces of rubber with short T2. All scans share the
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same nominal isotropic resolution, repetition time, and flip angle. (a) The gradient echo image is dominated by the oil signal while rubber is hardly visible. This is due to signal losses occurring before the echo is formed and acquired around the echo time TE. (b) In contrast, with zero echo time (ZTE)
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imaging, frequency encoding and acquisition start immediately after signal excitation, so that TE = 0. Hence, rubber appears in the image at full intensity. However, signal decay during the acquisition
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range Tk leads to image blurring, i.e., reduced actual spatial resolution for rubber. (c) Speeding up encoding by using a stronger gradient, thus creating a higher bandwidth BW, reduces the effects of signal decay and sharpens the edges of the rubber in the image. Hence, the two basic principles for
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encoding and acquisition in short-T2 MRI are 1) to start as early as possible and 2) to minimize the range Tk to avoid detrimental effects of signal decay. Fig. 4: High bandwidth as a requirement for short-T2 MRI. Imaging short-T2 samples at high spatial
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resolution requires both rapid encoding and large gradient (G) time-integrals. This can be achieved simultaneously only by using strong gradients, thus resulting in high resonance bandwidth. Fig. 5: Concepts and techniques for short-T2 MRI. The concepts follow the two principles of (1) an
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early start and (2) a small range of data acquisition, as illustrated in Fig. 3 and indicated by the red
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labels. (1) is achieved by A) using 3D imaging without slice selection, B) employing radial center-out gradient encoding, and C) performing signal excitation in the presence of the encoding gradient. For concept B, using either pure Cartesian phase or pure radial frequency encoding patterns favors one or the other principle. Combinations of the different excitation and encoding schemes lead to the four basic techniques for short-T2 MRI, namely constant time, single point, ultra-short TE, and zero TE
imaging, along with their intermediate relatives. CTI = constant time imaging, SPI = single point imaging, SPRITE = single-point ramped imaging with T1 enhancement, MP mapping = multipoint mapping [382], UTE = ultra-short echo time, BLAST = back-projection low angle shot, RUFIS = rotating ultra-fast imaging sequence, ZTE = zero echo time, WASPI = water- and fat-suppressed
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
proton projection imaging, SWIFT = sweep imaging with Fourier transformation, PETRA = pointwise encoding time reduction with radial acquisition, HYFI = hybrid filling, GM = gradient-modulated, RHE = ramped hybrid encoding. Fig. 6: Sequence diagrams of the four basic techniques for short-T2 MRI. They all employ 3D centerout k-space encoding from which one TR interval is depicted. (a) In constant time imaging (CTI), pure
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phase encoding is performed using a gradient of constant length and variable amplitude. One data
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point is acquired per excitation after identical TE, including k = 0. Thus, spatial resolution is not impaired by T2 decay but signal is reduced related to TE. (b) Shorter TE is achieved with single point
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imaging (SPI) where gradient ramps are excluded from encoding by performing both excitation and acquisition of high bandwidth while the gradient is at full strength. As indicated in the scheme by the
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grey bars, gradients need not be switched off between successive TR intervals, thus enabling nearsilent operation. (c) In contrast to the previous techniques, ultra-short echo time (UTE) imaging employs pure frequency encoding. Thus higher SNR efficiency is achieved but T2 decay may affect
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spatial resolution. After non-selective excitation, data acquisition is started concurrently with gradient ramping, the minimum TE being determined by the RF dead time. (d) In zero echo time (ZTE)
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imaging, high-bandwidth excitation is performed after ramping up the gradient. TE is actually zero because k = 0 coincides with the center of the RF pulse. Data acquisition starts only after the RF dead time, thus leading to a gap in central k-space. Equivalently to SPI, ZTE imaging can be performed
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with strongly reduced acoustic noise.
Fig. 7: Patterns in k-space of the four basic techniques for short-T2 MRI, schematically showing a central plane of 3D k-space for each case. The distance between neighboring k-space points according
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to the Nyquist criterion is given as the dwell time DW = 1/BW for the bandwidth BW in the FOV. (a) Both CTI and SPI are usually performed with Cartesian sampling. One k-space point is obtained per excitation after radial encoding of varying speed and direction. Thus, all points are equally T2-
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weighted but with a relatively long TE. (b) In UTE imaging, data are acquired along the radial
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trajectory. In central k-space (yellow), encoding speed increases while ramping up the gradient, leading to a non-uniform radial sampling density. (c) In ZTE imaging, full, constant encoding speed is achieved but data are missing in central k-space (red) due to finite RF dead time. Additional data points are acquired by radial oversampling. Fig. 8: Effect of gradient ramping on depiction behavior in UTE imaging. A 1D simulation is shown using different ramp durations associated with different gradient slew rates. All times and durations are normalized with respect to ' which is the acquisition range at constant gradient without ramping, i.e., as in the ZTE case shown in Fig. 4d. Using the parameters FOV 240 mm, spatial resolution 1 mm, and gradient strength 25 mT/m results in ' = 0.47 ms. A T2 value shorter than
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
' of 0.20 ms is assumed to emphasize the effects of signal decay. As a reference, the ideal case is shown where all data are obtained without any delay after signal generation, i.e., unaffected by T2. (a) Time of acquisition of the k-space positions of one radial readout. During the ramp, as indicated by the dashed lines, k-space speed increases until the target gradient strength is achieved. Compared to ', the UTE acquisition range is increased by half the ramping time. (b) Modulation transfer
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function (MTF) describing the apodization in k-space due to exponential T2 decay according to the
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delayed acquisition times shown in a). (c) Point-spread-function (PSF) obtained by symmetrically complementing, zero-filling, and Fourier-transforming the MTF in b). The displayed PSFs describe
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the distribution of signal from a pixel center (position 0) to other reconstructed pixels in the image and are depicted for the first 12 neighboring pixels. In the ideal case, the PSF is 1 at the location of origin
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and 0 at all other pixel positions. The Sinc shape reflects finite sampling up to kmax. For UTE encoding, the PSFs show a broadening of the main lobe and damped Sinc side lobes. (d) Normalizing
slew rates further emphasize this effect.
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the UTE PSFs to their maximum value reveals the associated loss of spatial resolution. Low gradient
Fig. 9: Impact of gradient fidelity on UTE image quality. (a) Ideal gradient shape with ramp. (b)
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Simulated image of a sphere obtained from data acquired with the ideal shape in a). (c) Distortion of gradient shape obtained for an exponential first-order, so-called gradient eddy current with time constant 0.05 '. (d) Image obtained when reconstructing data acquired with the distorted gradient
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in c) using the nominal, ideal trajectory. (e) Phase evolution from a zeroth-order field generated by an exponential, so-called B0 eddy current with time constant 0.5 '. (f) Image obtained from data with B0 eddy current in e) induced by the gradient coil encoding the horizontal direction. Correcting
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data for the B0-eddy-current-induced phase and reconstructing based on the k-space trajectory distorted by the gradient eddy current results in an image virtually identical to b). Fig. 10: Contributions to the initial RF dead time in pulsed MR. For symmetric RF pulses of constant
U
frequency, the focal point is located in the pulse center, so that only half of its duration contributes to
JO
the dead time. During T/R switching, the MR signal is corrupted by spikes and transients. After these RF events, fully filtered and uncorrupted data are only available after the lead-in time of the digital filter, corresponding to half of its length in the case of a symmetric filter with finite impulse response. The dwell time DW corresponds to the bandwidth in the target FOV according to Nyquist (see Fig. 7), whereas DWFOV is smaller and associated with radial oversampling. Fig. 11: Illustration of 1D algebraic reconstruction (AR) of a ZTE image. (a) Simulated ZTE k-space data obtained for a box-shaped object, exhibiting a gap of 3 Nyquist dwell times on the radius and two-fold oversampling. (b) Image obtained by FT of the data after filling the gap with zeroes, leading to low-frequency artifacts. (c) Image obtained by AR which is free of artifacts. (d) Data in k-space
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
obtained as FT of the reconstructed ZTE image, demonstrating recovery of data in the gap along with elimination of the oversampling. Fig. 12: Noise behavior in ZTE imaging. (a) 1D image reconstructed from pure noise time-domain data. Bad conditioning at a gap size of 3 Nyquist dwell times DW = 1/BW leads to noise enhancement and correlation in the image domain, which manifests as a low-frequency image distortion. Another
F
realization of the noise input data would lead to a different shape in the image. (b) Using as input the
O
same noise plus the signal from Fig. 11a results in a superposition of the two components also in the image. (c) Reconstruction of a 3D image from noise projections with gap = 3 DW as a). Due to high
O
inherent oversampling of 3D radial data in central k-space, noise enhancement is averaged out. (d) Only at larger gap = 4, further enhanced and correlated noise patterns are observed also in the 3D
PR
image.
Fig. 13: Background signal in ZTE imaging. Images of a human head were acquired at 7 T using a
E-
T/R birdcage RF coil. All kinds of signals were detected, including signals from very short-T2 materials of the RF coil and cables, the head support, and the patient table. (a) With a small dead-time
PR
gap and a FOV adapted to the head size, ordinary aliasing of radial data is manifested at the image center and borders. (b) A FOV twice as large contains most of the background signal and aliasing is largely eliminated. However, scan time is four-fold, and bandwidth BW and the gap size given in Nyquist dwell times DW = 1/BW are doubled. (c) For a slightly larger gap, worse conditioning
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amplifies the aliased out-of-band signal. (d) Choosing a larger FOV does not help in this case as the increased gap of 3 DW also leads to amplification of residual aliased signal.
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Fig. 14: RF excitation for ZTE imaging for a given signal bandwidth BW in the FOV. (a) Time course of short and long block-shaped pulses and a modulated pulse. Durations are given in Nyquist dwell times DW = 1/BW. The frequency sweep of the modulated pulse approximately covers BW. (b) Spectral amplitudes exhibiting high uniformity for the short block (= hard) pulse and strong variations
U
for the longer block pulse. The modulated pulse achieves about three times the amplitude of the hard
JO
pulse for a comparable uniformity. (c) ZTE image of a sphere using the short block pulse, showing uniform excitation. (d) A longer block pulse of duration 2.5 DW results in strong variations in image intensity. (e) A pulse of equal duration but with amplitude and frequency modulation provides uniformity comparable to the short block pulse. Fig. 15: RF excitation using a long modulated pulse as done in the SWIFT technique. To avoid missing signals generated and encoded during earlier parts of the pulse, excitation and acquisition are performed quasi-simultaneously. This is achieved by periodically alternating between the two operation modes. Thus both excitation and acquisition are run at a reduced duty cycle which is further
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
decreased by switching dead times. To also fully encode signals created at the end of the pulse, the encoding gradient and acquisition need to be continued after excitation [107]. Fig. 16: Patterns in k-space of two techniques filling the dead-time gap in ZTE imaging. (a) In PETRA, the gap area is covered by additional Cartesian SPI acquisitions. (b) In WASPI, additional radial projections are acquired at reduced gradient strength to limit the gap size to half a Nyquist
F
distance. To limit extra scan time, the angular density is chosen to meet the Nyquist criterion only at
O
the interface between the two parts of the data.
Fig. 17: Image fidelity of ZTE imaging and related techniques, illustrated in analogy to Fig. 8. (a) In
O
SPI, the time of acquisition is constant as in the ideal case but delayed according to the largest
PR
encoding. ZTE encoding starts at 0, increases linearly, and ends at the SPI delay. Note that with the present gap size of 50 μs = 13 DW, the ZTE case is not actually possible due to ill-conditioning and is only shown to illustrate the purely exponential decay. With PETRA, the additional SPI data in the gap
E-
are acquired after the dead time, leading to a reduced acquisition range Tk. In WASPI, a discontinuity is created by the different gradient strengths. (b) T2 decay illustrated by the MTF shows no effect for
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SPI, exponential weighting for ZTE acquisition, a flattened initial part for PETRA, and a discontinuous behavior for WASPI. (c) The PSF shows that at the original pixel location the strongest signal reduction is experienced by SPI, whereas ZTE, WASPI, and PETRA exhibit a comparable decrease. (d) The normalized PSF shows ideal resolution for SPI and an increase of the main lobe
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width for the three ZTE relatives. For both ZTE and PETRA, a resolution loss is observed from the distribution of signals to neighboring pixels. With WASPI, this effect is stronger and reaches further
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out, thus leading to side lobes.
Fig. 18: Generating contrast in ZTE imaging by magnetization preparation demonstrated on an excised lamb joint. (a) Non-prepared ZTE image showing primarily proton density and some T1 contrast. (b) In gradient echo imaging, signals from short-T2 tissues are reduced, in particular from
U
cortical bone. (c) In ZTE imaging with fat suppression, both short- and long-T2 water signals remain.
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(d) Selectively suppressing long-T2 water in the ZTE sequence removes signals from cartilage and
muscle, and fat signal dominates the image. (e) Applying suppression of both fat and long-T2 water results in positive contrast for short-T2 tissues, showing primarily trabecular and cortical bone. Remaining bright fat signal (right) indicates imperfect suppression due to local off-resonance. (f) In the CT reference image, strong similarities with the MRI data in e) are observed. Reproduced with permission from Ref. [153]. Fig. 19: Off-resonance effects in radial imaging, demonstrated for the specific case of chemical shifts in the presence of both water and fat. (a) Simulated images with a pixel bandwidth (PBW) corresponding to the frequency difference between water and fat. Setting the demodulation frequency
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
f to the water resonance results in sharp edges of the water compartment and blurred depiction of the fat parts. At interfaces of the two compartments, destructive and constructive interference lead to edge enhancement (arrow). Demodulation at the fat frequency changes the order of low and high signal intensities at interfaces whereas at the intermediate frequency, edges are not enhanced but blurred. (b) In-vivo ZTE data of a human knee acquired at 7 T with PBW = 870 Hz, corresponding to 0.87 times
F
the water-to-fat-shift. At different demodulation frequencies, the equivalent effects to the simulations
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in a) are observed, as indicated by arrows.
Fig. 20: Resolution enhancement by exponential T2 filtering in simulated radial MRI with center-out
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encoding. (a) Two ellipse-shaped objects without T2 filtering. For a relatively long T2 equal to the acquisition range Tk, the object is sharply depicted. For a significantly shorter T2, blurring is observed.
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(b) Moderate T2 filtering enhances resolution of the short-T2 object but leads to edge artifacts for the object with longer T2. (c) With full compensation of T2 decay, the short-T2 object is sharply depicted.
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However, strong artifacts are observed for the long-T2 object and noise is enhanced.
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Fig. 21: Selected short-T2 MRI applications. All images were obtained using the 3D ZTE technique. (a) MSK: Human knee acquired at 7 T with isotropic spatial resolution of 0.75 mm in 4 m 17 s. Reproduced with permission from Ref. [205]. (b) Pulmonary MRI: Mouse lung acquired at 4.7 T with a resolution of 0.31 mm in 4 m 30 s using respiratory gating. Reproduced with permission from Ref.
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[99]. (c) Dental MRI: Excised human tooth acquired at 11.7 T with a resolution 0.13 × 0.13 × 0.18
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mm3 in 21 m 18 s. Reproduced with permission from Ref. [310].
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Table 1: Principles for short-T2 MRI. = time of acquisition, TE = echo time, SNR = signal-tonoise ratio, = range of times of acquisition.
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Principle for data acquisition 1. Early start
2. Small range
Characterized by
Affects
!" , TE
SNR
Spatial resolution
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Table 2: Examples of short-T2 tissues and materials. Relaxation times were derived and selected from single- or multi-component analysis. An asterisk * indicates a T2* value as opposed to T2. Unless stated otherwise, the values associated with tissues refer to 1H signal in water molecules. B0 [T]
Origin
T2 or T2* [μs]
References
Cortical bone
0.47
Porcine, bovine, specimen
400
[347]
(bone water)
0.64
Human, specimen
400
[348]
1.5
Human
420 – 500 *
Human, specimen
401 *
Human
260 – 786 *
Human, specimen
237 – 576 *
[224, 233, 349]
Bovine, specimen
290 – 368 *
[175, 227, 350-352]
4.7
Human, specimen Human, specimen
7.0
O
O
290 *
[354]
431 *
[168]
290 *
[236]
Human, specimen
302 *
[349]
Human, specimen
540, 400 *
[243]
Porcine, specimen
597 *
[258]
Human
176 – 207 *
[252, 355]
Bovine, specimen
190 *
[356]
Human
160 *
[249]
Human
178 *
[256]
Lamb, specimen
189 *
[250]
Synthesized
448
[248]
7.0
Chicken, specimen
675
[248]
9.4
Lamb, specimen
119 *
[250]
Rabbit, specimen
90 *
[22]
0.47
Porcine, specimen
650 – 3800
[25]
1.5
Human
1200 – 2000 *
[150, 259]
3.0
Human
5200 – 6000
[161]
Human
600 – 2320 *
[31, 54, 173, 227, 260, 357, 358]
Human, specimen
1180 – 2180 *
Bovine, specimen
1280 *
Human
340
AL
1.5
PR
Bovine, specimen
3.0
R N U JO
[353, 354]
[349]
5.6
Tendons
400 – 416
368 *
11.7
(bone matrix, hydroxyapatite 31P)
[53, 166, 172, 231]
Human, specimen
9.4
Cortical bone
[349]
E-
Mouse
[232]
PR
3.0
F
Tissue/material
7.0
[359, 360] [350] [361]
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Collagen protein
Dentine
[51]
3.0
Human
5800 – 9300
[161]
Human
3260 *
[173]
Caprine, specimen
1530 *
[350]
1.5
Human
3000 *
[150]
3.0
Human
9700 – 11400
[161]
Human
4190 *
Human, specimen
1800 – 2100 *
1.5
Human
5000 – 10000
3.0
Human
1340 *
Human, specimen
480 – 990 *
9.4
Bovine, specimen
0.47
O [51]
O
[173]
[264, 350]
Porcine, bovine, specimen
12 *
[347]
0.7
Bovine, specimen
20 *
[363]
1.4
Rat tendon, specimen
90
[32]
4.7
Human, specimen
57, 12 *
[353]
9.4
Rabbit, specimen
20 – 40 *
[228]
Human, specimen
38000
[364]
Human, specimen
12, 65, 200, 1000 *
[364]
Human
324
[307]
Human, specimen
12, 65, 230 *
[365]
0.9
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0.9 1.5
Human
150 – 350 *
[366]
3.0
Human
358 *
[320]
Human
420 *
[147]
Human
500 – 700 *
[324]
Human
260 *
[367]
Human
336 *
[368]
Human, MS specimen
216 *
[320]
Ovine, specimen
200 – 300 *
[369]
Ovine, specimen
150 – 400 *
[370]
Bovine, specimen
114 *
[320]
4.7
Rat, specimen
50 – 1000 *
[318]
7.0
Human
200 – 300 *
[324]
9.4
Human, specimen
38 – 57
[371]
Rat, bovine, specimen
8 – 1000 *
[35]
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[227, 350]
[362]
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Myelin lipids
[173]
2300
3.0 Dental enamel
F
4000 – 10000
PR
Cartilage
Human
E-
Meniscus
1.5
PR
Ligaments
JOURNAL PRE-PROOF Weiger et al., Short-T2 MRI
Human
4000 *
[372]
0.68
Rat, specimen
2340 *
[373]
1.5
Human
1100 – 2110 *
[287, 293, 374-377]
2.0
Rat, specimen
830 *
[373]
3.0
Human
460 – 820 *
[284, 293]
Mouse
910 – 1000 *
[288, 292]
4.7
Mouse
460 – 500 *
7.0
Mouse
395 *
7.9
Mouse
770 *
Liver
3.0
Human
300 – 450 *
Rubber
2.5
Polyvinylchloride (PVC)
2.4
Acrylonitrile butadiene styrene (ABS)
2.5
R N U JO
O
[162] [167]
O
[379, 380] [381]
180 – 1350 *
[147, 172, 223]
38 *
[340]
98 *
[340]
E-
2.4
PR
Polyethylene (PE)
AL
2.4
[285, 378]
300 – 1300 *
3.0 Polymethylmethacrylate (PMMA, acrylic glass)
F
0.5
PR
Lung
442 *
[340]
760 *
[381]
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a
Molecular tumbling or crystal rotation Dipolar field
B0 θ
F
θ
B0 + ΔB(θ)
Spectrum
E-
Liquid
PR
b
O
O
Molecule
Single crystal
θ = 20°
PR
Liquid with relaxation
θ = 90°
Powder
AL
θ = 55°
f
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0
θ = 0°
Free induction decay
JO
U
c
0
Fig. 1
t
Fig. 2
ΔB
Amplitude
Interaction (e.g. dipolar coupling)
τc
t [s]
τc = 0.05 ns
τc = 0.4 ns
W0
W2
AL W1
W 1 ~ J(ω0)
W 2 ~ J(2ω0)
Transition probabilities
W 0 ~ J(0)
FT
[s]
Liquids
F
O
τc [s]
τc = 1/ω0
Relaxation
ω [rad/s]
T2
T1
Spectral density
0 ω0 2ω0
J
O
PR
E-
Weights
PR τ [s]
Autocorrelation function
Energy diagram spin pair
AC
R N
U
JO
Random field
Solids
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JO
U
Gradient echo
t
TE
AQ
Signal loss
Rubber T2 ≈ 0.4 ms
Oil T2 ≈ 30 ms
b ZTE
AQ
1 Early start
TE = 0
Tk
c ZTE
F
2 Small range
O
AQ
Tk
Tk = 0.32 ms, BW = 200 kHz
O
PR Signal decay
E-
PR
Tk = 1.28 ms, BW = 50 kHz
AL
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TE = 1.32 ms, BW = 100 kHz
Fig. 3
G
RF
Sig
a
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Fig. 4
+
High bandwidth
F
O
O
PR
Large G-integral
High resolution
E-
PR
Strong G
AL
Rapid encoding
Short T2
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U
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Fig. 5
1
Radial center-out encoding
B
1
Frequency encoding Radial
Intermediate
2
Phase encoding Cartesian
C
A
AL
Ultra-short TE UTE
Constant Time Imaging CTI
GM-SWIFT
MP mapping WASPI PETRA HYFI
Single Point Imaging SPI, SPRITE
RF during G
3D
2
1
1
Zero TE BLAST, RUFIS, ZTE SWIFT
F
O
O
PR
RHE, GM-PETRA
E-
Intermediate
No slice selection
PR
RF before G
R N
U
Concept
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TE
AQ
Tk
PR
TE = 0
F
AQ
Tk
TE
O
O
Dead time
d ZTE
b SPI
E-
PR
k=0
AQ
Tk = 0
AL
R N
TE
U
JO
Ramp
c UTE
Fig. 6
G
RF
G
RF
a CTI
AQ
Tk = 0
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gradient encoding
b
k-space points:
DW
Nyquist
k0
F
O
missed
Dead time gap
ZTE
oversampled
O
c
PR acquired
E-
PR
UTE
Gradient ramp
AL
R N
U
JO
CTI / SPI
Equal, long TE
Fig. 7
a
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Fig. 8
b
ramp
AL
R N
U
JO
a
Tk
d
F
O
O
PR
E-
PR
plateau
c
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Fig. 9
AL
R N
U
JO
F
O
O
PR
E-
PR
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Fig. 10
RF pulse
F
DW
O
O
PR
E-
PR
T/R switching Filter lead-in
AL
Dead time
R N
U
JO
DWov
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Fig. 11
d
a
Dead time gap
IFT
FT
c
b
AL
R N
U
JO
k-Space
F
O
O
PR
E-
PR
Image
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c
Fig. 12
3D
1D
a
gap = 3
gap = 3
Noise
d
gap = 4
gap = 3
AL
Signal + noise
R N
b
U
JO
F
O
O
PR
E-
PR
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Fig. 13
FOV 530 mm BW 500 kHz DW 2 μs
FOV 265 mm BW 250 kHz DW 4 μs
Gap 2
b
Gap 1
a
AL
R N
Dead time 4 μs
U
JO
F
O
O
PR Gap 3
E-
PR d
Gap 1.5
c
Dead time 6 μs
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Fig. 14
b
a
-1
-0.5
0
BW = FOV
0
0.5
f/BW
1 t/DW
d
c
2.5 DW
0.4 DW
e 2.5 DW
F
O
O
PR
E-
PR
BW
AL
Block short Block long Mod. amp. Mod. freq.
R N
U
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Freq
Fig. 15
G
AQ
RF
Amp
U
RF AQ Dead interval interval time
F
O
O
PR
E-
PR
Encoding time
AL
R N
Pulse duration
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Fig. 16
a
SPI
b
F
O
O
PR
E-
Lower gradient
PR
WASPI
AL
R N
U
JO
PETRA
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Fig. 17
b
gap
AL
R N
U
JO
a
d
F
O
O
PR
E-
PR
Tk
c
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fat
Gradient echo
d
c
Water sup
f
CT
F
O
O
PR
E-
e Fat + water sup
PR
cartilage
AL
Fat sup
R N
U
trab. cort. bone bone
ZTE
muscle
Fig. 18
b
a
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fat water
F
O
f = ffat
O
PR
E-
PR
f = (fwater+ffat)/2
AL
R N
U
JO
PBW = fwater - ffat
Fig. 19
b
a
f = fwater
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b
c
F
T2,filter = T2,short
O
O
PR
E-
PR
T2,filter = 2 T2,short
AL
R N
U
T2,short = 0.2 Tk
No T2 filtering
T2,long = Tk
Fig. 20
a
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Fig. 21
a
patellar tendon
b
AL
R N
U
JO
F
O
O
pulp
PR
E-
PR
caries lesion
c
bone
dentine
enamel
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JO
U
R N
AL
PR
E-
PR
O
O
F
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JOURNAL PRE-PROOF Short-T2 MRI: Principles and recent advances Highlights Overarching concepts structure the variety of short-T2 MRI techniques Relaxation in a nutshell
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Detailed description of ZTE imaging and its variants
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Dedicated hardware requirements for all scanner components
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E-
PR
O
Comprehensive literature overview including T2 and T2* values of potential imaging targets