Imaging of short and ultrashort T2 and T2 * tissues using clinical MRI systems

Graeme M Bydder^*

Department of Radiology, University of California, San Diego, 200 West Arbor Drive, San Diego, CA 92103–8226, USA

Corresponding Author:

Graeme M Bydder
Department of Radiology
University of California, San Diego
200 West Arbor Drive, San Diego
CA 92103–8226, USA
Tel: +1 619 471 0506
Fax: +1 619 471 0503
E-mail: gbydder@ucsd.edu

Abstract

There are now a variety of new techniques available to detect signal from tissues with short or ultrashort T2s and T2*s. There are also many methods of developing image contrast between tissues and fluids in the short T2 or T2* range, which can provide visualization of anatomy that has not previously been seen. Particular methods have been developed to target susceptibility effects, and allow accurate quantitation by compensating for the anatomical distortion produced by these effects. Specific methods have been developed to image the effects of magnetic iron oxide particles with positive contrast and to correct for the loss of signal and image distortion near to metal caused by gross susceptibility effects. These methods are likely to provide interesting options and increase the range of applications of MRI.

Keywords

short T₂ tissue components ▪ susceptibility ▪ ultrashort echo time

During the first year of clinical MRI, only steady state free precession, T₁-weighted and proton density-weighted clinical images were available [1–3]. Heavily T₂-weighted spin echo (SE) sequences arrived suddenly in early 1982 and transformed the practice of magnetic resonance (MR) [4–6]. Images obtained with these sequences detected intermediate or long T₂ relaxation components in tissue. Even with the subsequent development of new classes of sequences, such as fast SE, clinical diffusion-weighted imaging and fluid attenuated inversion recovery, the detection of signal from intermediate and long T₂ relaxation components remains the dominant form of MRI for the diagnosis of parenchymal disease in the brain and much of the rest of the body.

However, even when clinical MRI began, very short mean T₂ relaxation components were recognized in the cortical bone by Smith et al. [7] and Edelstein et al. [8]. This tissue showed no MR signal. The lack of signal was useful in providing a low-signal background against which abnormalities in cortical bone with sufficiently long mean T₂s to result in detectable signal could be recognized, but the absence of signal from normal cortical bone meant that there was no possibility of measuring normal values of mobile proton density (ρ_m), T₁ or T₂. Nor was it possible to study normal perfusion, and there was no opportunity for active contrast manipulation, little or no distinction between adjacent short T₂ tissues and no means of visualizing normal contrast enhancement. As a result, the study of cortical bone and other MR ‘invisible’ short T₂ tissues, such as tendons, ligaments and menisci, has been far more limited than that of tissues and organs, such as brain, liver and muscle, where tissue mean T₂s are longer and MR signal from them is readily detectable with conventional clinical sequences. However, even these longer T₂ tissues contain significant proportions (e.g., 5–30%) of invisible or undetectable short T₂ relaxation components when they are imaged with conventional approaches.

To image short or ultrashort T₂ tissues that produce no detectable signal with conventional sequences, indirect methods have been used in which signal is obtained from surrounding or associated longer T₂ tissues. When the low- or zero-signal tissue is surrounded by longer T₂ tissue, signal from this tissue can be used to define the boundaries of the zero-signal tissue. It is also possible to characterize some short T₂ tissues by observing the impact that their difference in susceptibility from that of the surrounding longer T₂ tissues has on the signal obtained from the longer T₂ tissue. For example, some features of trabecular bone can be inferred by the effect this tissue has on the MR signal of adjacent red or yellow bone marrow [9]. A third indirect method is possible when short and long T₂ relaxation components are associated, and undergo magnetization exchange. The effect of saturation of the invisible short T₂ components on this exchange can be observed on the detectable longer T₂ components [10] and, thus, inferences can be made regarding the short T₂ tissue and/or the exchange between the shorter and longer T₂ components.

An alternative to using conventional sequences to study short T₂ tissues in this way is to employ methods that directly detect signal from them. These usually involve the use of short echo time (TE) or ultrashort TE (UTE) sequences to detect MR signals before they have decayed to zero. There are now a variety of sequences of this type available in the clinical domain.

While T₂ is a property of tissue that reflects dipolar and other nuclear interactions, frequently the effects seen with MRI are described more accurately by the observed T₂ or T₂*. This includes effects such as intravoxel dephasing due to B₀ field inhomogeneity, tissue susceptibility differences and chemical shift. Tissue susceptibility effects reflect the fact that solid tissues, such as bone, are generally more diamagnetic than soft tissues, and that some tissues and fluids may be paramagnetic. The effects of some of these differences can be partly or almost wholly reversed by the use of SE sequences.

In some situations, T₂* effects may dominate and it is useful to recognize several different approaches to imaging short T₂/T₂* components in relation to underlying susceptibility differences:

▪ The first approach essentially sees the problem as imaging of short or ultrashort T₂ components and the basic approach is to use a short TE or UTE to acquire and encode MR signals before they decay to a low level. This may be appropriate in situations where there are only minor susceptibility differences present.

▪ The second is susceptibility-weighted imaging, where magnitude and phase data are used to recognize a loss of signal from tissue caused by susceptibility effects. It can be direct and/or indirect (where T₂* become too short to detect) and is qualitative.

▪ Quantitative susceptibility imaging is the third approach. This technique recognizes the fact that susceptibility differences affect the spatial encoding of MR signals and endeavors to correct this and to calculate values of T₂* that accurately reflect T₂ and susceptibility effects.

▪ Positive-contrast and white-marker imaging techniques address the specific problem of imaging the effects of magnetic iron oxide particles (MIOPs), which shorten T₂ and produce local disturbances of the magnetic field. The aim is to detect the presence of particles with a positive signal and at least, in part, address the problem of field distortion to achieve credible recognition and quantification of the concentration of MIOPs.

▪ The fifth group of techniques is targeted at imaging in the presence of metal. Metals can produce very large susceptibility effects with loss of signal due to T₂* effects and gross image distortion. The primary objective in this situation is to deal with the image distortion and restore image integrity to a sufficient degree for the images to be clinically useful.

There is some overlap between these approaches. In some situations it may be appropriate to ignore the effects of susceptibility differences in producing image distortion and regard the problem as one of detecting short T₂ signals, whereas in other situations image distortion due to susceptibility is the primary problem that needs to be addressed. Over the previous year, there has been considerable interest in these approaches with solutions now appearing to some problems that were previously intractable for many years.

Tissue properties

The tissues of the human body can be divided into those that are visible, in the sense that they provide detectable signal with clinical MR systems, and those that are ‘invisible’ because their mean T₂s or T₂*s are too short to provide a detectable signal. All tissues have multicomponent T₂s. This means that they contain a mixture of short and long T₂ components. The invisible tissues have a majority of short T₂ components and a minority of long T₂ components. The latter components typically do not provide enough signal to be detectable in comparison to image noise levels. The invisible tissues of the body, such as the brain, liver and muscle, have a majority of long T₂ components, which produce the signals seen with conventional techniques. They also have a minority of short T₂ components that do not contribute significantly to the detectable signal.

There is no agreement as to what constitutes a short TE and what is an UTE, and there is an argument regarding how TE should be measured for tissues with short T₂s [11], but for simplicity, a short TE is taken to be less than 10 ms and a UTE less than 1 ms. It is also possible to define short T₂/T₂* as less than 10 ms and ultrashort as less than 1 ms. This reflects the fact that with older systems and SE sequences, tissues with a T₂ or T₂* less than 10 ms produced little or no signal and were invisible. With more recent systems and gradient-echo sequences the cut-off is closer to 1 ms.

Within the invisible group of tissues (mean T₂ < 10 ms) it is possible to differentiate a first group, including tendons, ligaments and menisci, with short mean T₂s of approximately 1–10 ms, a second group, including cortical bone and dentine, with ultrashort mean T₂s of 0.1–1 ms. There is also a third group, including dental enamel, protons in membranes, large molecules and crystalline bone, with a mean T₂s of less than 0.1 ms.

An important factor in this context is the magic-angle effect [12,13] since it can greatly increase the T₂ of short T₂ tissues, such as tendons, ligaments and menisci. When the orientation of tissues that contain highly ordered collagen is changed, their T₂ varies from a minimum at θ = 0°, where dipolar interactions are greatest, to a maximum where 3cos²θ – 1 ≈ 0 and θ = 55°. q is the orientation of the fibers to B₀. The increase can be large, for example, from 0.6 to 21 ms [12] or from 7 to 23 ms [13] in the Achilles tendon.

A recently described phenomenon is directional susceptibility in tendons, whereby their bulk magnetic susceptibility varies with orientation to B₀, with signals at the water end of the proton spectrum when fibers are parallel to B₀ and at the fat end of the spectrum (lower frequency) when fibers are perpendicular to B₀ [14]. The difference is relatively large (of the order of three parts per million [ppm]).

The ρ_m of tissues also varies markedly, with bone having a ρ_m of 15–20% and semi-solid tissues, such as tendons and ligaments, having values of 60–70%. ρ_m is generally a more important factor in generating contrast with short T₂ tissues than it is with longer T₂ tissues. The low ρ_m for bone places a limit on the maximum signal that can be obtained from it.

The mean T₁s of some tissues with a majority of short T₂ components are short, with cortical bone having a particularly short T₁, in fact, less than that of fat [15]. The relative differences in mean T₂ or T₂* between normal and abnormal tissue are generally much greater than those in mean T₁.

Relative to air, soft tissues generally show a susceptibility difference of approximately -9 ppm, and bone and calcified tissue of approximately -11 ppm. By comparison, the principal peak of fat resonates at approximately -12 ppm. Paramagnetic materials may show small positive chemical shifts and superparamagnetic materials show greater positive shifts again. Metals including, for example, titanium and some types of stainless steel, may show large positive shifts of 10–100s ppm (or more). These changes in field may be considerably greater than those used by applied machine gradients to encode MR signals and may therefore cause image distortions.

Acquisition methods for short T₂/T₂* components of tissue

Some of the techniques currently being used to detect signal directly from tissues on clinical systems have been used in material science and tissue studies using small-bore high-field spectrometers for many years. Methods now in use on clinical systems are summarized in Table 1. The prototype sequence for imaging short T₂ tissues is single-point imaging where a single point in k-space is acquired with an UTE. This is typically used with 3D phase encoding, which unfortunately makes the technique very time consuming [16].

It is possible to acquire several points at a time, which makes the sequences more time efficient but results in longer TEs for the additional points [17]. There are also free induction decaybased techniques where a radial line of k-space is acquired from the center out [18]. This can be coupled with long T₂ water- and fat-suppression to selectively image short T₂ components [19]. Other trajectories in k-space are possible, including a stack of spirals [20].

A particularly innovative method of imaging short T₂ components is to divide the excitation pulse into subpulses and acquire data after each of these pulses. The acquired data need to be deconvolved with the excitation pulse, but the end result is a much more time efficient acquisition than with typical 3D acquisitions [21–24]. Other techniques that have only been used in the preclinical phase include methods in which radiofrequency absorption, rather than signal detection, is assessed [25]. The methods borrow from older forms of spectroscopy and electron spin resonance, where electronic T₂s are extremely short and may be of the order of a microsecond.

Magnetization preparation, contrast mechanisms & signal-suppression techniques

Traditional contrast mechanisms exploiting differences in ρ_m, chemical shift and other tissue properties can be used in ways that are well known from conventional imaging.

There are also numerous new contrast mechanisms, or old contrast mechanisms operating in new ways, that are of interest in imaging short/ultrashort T₂/T₂* components in tissue. Some of these are listed in Table 2. They are typically used in conjunction with the acquisition techniques detailed in the previous section. These provide a wide range of possible ways of effecting magnetization. For example, 90 and 180°, fat saturation and magnetization transfer pulses can all be used to suppress unwanted long T₂ signals and to produce T₂ contrast in the short T₂ range. There are also relatively new potential mechanisms involving double quantum filters [26] and a reduction in dipolar coupling [27,28]. These techniques are usually applied in conjunction with one of the acquisition methods described in the previous section.

Susceptibility-weighted imaging

Susceptibility-weighted imaging has been in use for a considerable time. It usually exploits reductions in T₂* to develop contrast, and imaging may utilize both magnitude and phase data [29,30]. The T₂* may be so short that this, in effect, becomes an indirect form of imaging utilizing the reduction in signal of adjacent longer T₂ components. The applicability of the technique and related methods can be expanded by utilizing forms of data collection with TEs or UTEs that can detect signal from very short T₂* components [31].

Quantitative susceptibility imaging

Quantitative methods of imaging susceptibility changes need to account for errors in spatial encoding, which may require solutions to a complex inverse problem [32,33]. To date, it has mainly been applied to brain imaging.

Positive-contrast & white-marker imaging

These forms of imaging have been used to describe the particular situation with MIOPs that may not only reduce T₂ and T₂*, but also produce local field distortions. A variety of different methods are available. It is possible to selectively excite only off-resonance spins. It is also possible to apply an additional gradient so that only the magnetization of spins in regions affected by MIOPs are refocused. The inhomogeneities from the particles induce echo shifts and these can be used to calculate and correct for the field distortion. The images reflect both tissue MIOP concentration and deviations of the local magnetic field produced by the particles [34–38].

Imaging in the presence of metal

When forms of metal are implanted in the body an extreme situation may arise in which there is very marked T₂* shortening, but the image distortion is so great that images of regions adjacent to the metal are uninterpretable. This has been a longstanding problem. In the past, a variety of solutions have been proposed, but these have had relatively little clinical impact. The recent development of multiacquisition variableresonance image combination (MAVRIC) [39] and slice encoding for metal artifact correction (SEMAC) [40] has resulted in a remarkable degree of restoration of images that are grossly degraded by a metallic artifact when imaged using conventional approaches. With MAVRIC, irradiation at a range of different off-resonance frequencies is used to detect signals whose resonant frequency has been shifted by metal. With SEMAC, phase encoding is used during slice selection to reallocate signals that are improperly located by the slice selection process. View angle tilting is also used with this technique to correct for errors with in-plane spatial encoding [41].

Imaging of boundaries involving short T₂/T₂* tissues

Structures of interest in the short T₂ range include thin layers such as those in entheses, periosteum and the deep layers of articular cartilage where there are short T₂ tissues, susceptibility effects between the soft (or semi-solid) tissues and bone, as well as partial volume effects between these tissues over curved surfaces. In this situation, high-resolution 3D-isotropic UTE imaging often has a distinct advantage since it can detect short T₂/T₂* signals as well as reduce the impact of susceptibility differences and partial volume effects. Imaging of ordered fibrous structures, such as tendons and ligaments, include some of the previously discussed issues but, in addition, a loss of contrast of the fiber structure or a blurred appearance may arise from obliquity of the fibers relative to the imaging slice. This Filler effect may simulate changes due to disease. There are also distinctive artifacts at boundaries from chemical shift effects, including those associated with radial acquisitions.

Clinical applications

There are now 2D and 3D UTE sequences available with imaging times of 5–6 min and clinically acceptable spatial resolution. In general the difficulty of acquiring short/ultrashort T₂/T₂* signals means that invisible tissues are imaged at lower spatial resolutions, but with signal levels and contrast that are not attainable with conventional techniques. A balance is necessary to obtain novel qualitative and/or quantitative information at spatial resolutions that show anatomical features with acceptable clarity.

Cortical bone

Cortical bone can be imaged with high signal [15]. The T₂ is approximately 0.4 ms and T₁ 250–350 ms at 1.5 T, which is shorter than typical values for fat. The mobile proton density is approximately 15–20%. These data can be used both for quantitative [42] and qualitative studies (Figure 1).

Tendons, ligaments & entheses

In conventional sequences, the signal from tendons, ligaments and entheses is very low or zero. Entheses are the attachment sites of tendons, ligaments and capsules to bone. They typically contain calcified and uncalcified fibrocartilage (which both have short T₂s). These tissues have a major role in dispersing mechanical stress at the junction between flexible tendons or ligaments and rigid bone.

Tendons and ligaments contain endotenon and endoligament, which have longer T₂s than the fibrous components (although they are still in the short T₂ range) and less magic-angle effect. Uncalcified fibrocartilage has a longer T₂ than the tensile components of tendons as well as an increase in T₂ caused by the magic-angle effect, but this may be present over a wider range of angles reflecting the more dispersed arrangement of the fibers within it. Magic-angle effects may result in a longer T₂ adjacent to bone from fibers that change in direction, from parallel to the bone surface to perpendicular to it, as they insert into bone. Tendons and ligaments can readily be seen with UTE sequences and entheses have been studied in detail [43,44]. Offresonance fat-suppression pulses reduce the signal from short T₂ fibers (which have a broad line width) more than endotenon or enthesis fibrocartilage (which have longer T₂s and narrower line widths) and this can be an effective contrast mechanism. Inversion pulses may be used to selectively invert and null enthesis fibrocartilage (exploiting its longer T₂), and so, visualize this tissue with high contrast. It is also possible to visualize oblique and transverse fibers in tendons using a combination of fatsuppressed UTE sequences to reduce short T₂ tissue water components and magic-angle imaging to lengthen the T₂ of the fibers at particular angles to B₀ (Figure 2).

Entheses are selectively involved in the seronegative spondyloarthropathies, such as ankylosing spondylitis and psoriatic arthropathy. The differential diagnosis is of a loss or reduction in fascicular pattern and includes normal sesamoid fibrocartilage, partial volume effects with a loss of fascicular pattern due to the Filler effect, magic-angle effects and disease.

Meniscii of the knee

The central region of the adult meniscus has no blood supply (the white zone) while the more peripheral region (the red zone) has a blood supply. Healing of tears in the white zone is generally unsatisfactory and the preferred surgical strategy is usually resection of the torn tissue. Suture and repair is more successful in the red zone. Distinction between the two zones has not previously been possible with MRI using conventional sequences, despite repeated attempts [45]. However, using UTE sequence and gadolinium-based contrast enhancement, the two zones can be distinguished [46] and provide a basis for surgical planning.

Anatomical descriptions of the meniscus include circumferential, radial, lamella, vertical and meshwork fiber groups. With conventional imaging, some radial fibers may be distinguishable from the majority of circumferential fibers, but with UTE and magic-angle imaging each of these fiber groups can be identified (Figure 3). It is also possible to distinguish the internal structure of the meniscus from that of the root ligaments, and the more central cartilaginous region from the more fibrous peripheral region of the meniscus.

The fiber structure provides a basis for understanding the biomechanics of the knee and the various patterns of tear in the meniscus. It also helps in distinguishing magic-angle effects within fiber groups from degenerative changes.

Temporomandibular joint disc

This demonstrated some of the characteristics of the meniscus of the knee. Fiber structure can be seen. Lamella, circumferential antero–posterior and superior–inferior fibers are identifiable.

Articular cartilage

Articular cartilage has a range of T₂s from approximately 1 to 30–40 ms from deep to superficial. With conventional imaging, the deep radial and calcified layers as well as the adjacent subchondrial bone are invisible. With UTE imaging, signal is detectable from the deeper layers of cartilage, allowing more superficial cartilage and subchondral bone to be distinguished [47]. This provides a basis for study of the junction between cartilage and bone, which may be of importance in the pathogenesis of osteoarthritis. Complex magic-angle effects are seen, owing to the fibrous architecture of articular cartilage.

In disease there may be both a loss of signal from the deep layer and an increased extent of the short T₂ associated with deep layers. There is electron microscope evidence of thinning of the deep layers in osteoarthritis but preservation in osteomalacia.

Spine

Imaging of the spine includes many visible tissues, therefore, to date, attention has focused on invisible structures such as entheses, the end plate of the disc, and short T₂ components in the intervertebral discs and red bone marrow. Fibrocartilage has also been demonstrated in the functional entheses of the transverse ligament of the atlas and the alar ligament. The dorsal capsule of the facet joints of the lumbar spine are also subject to cartilaginous metaplasia. Evidence of iron deposition can be seen in intervertebral discs in thalassemia [48].

Brain

There are significant short T₂ components in many tissues of the body with longer mean T₂s, including the brain, liver and muscle. These components can be specifically detected using UTE and other acquisition methods coupled with techniques that suppress long T₂ signals [49,50].

Liver

The liver contains a relatively high proportion of short T₂ components. The T₂*s of these may be prolonged in fibrosis [51]. The fibrosis in this situation is often of a relatively open structure and includes free water.

Pelvis

Ultrashort echo time sequences have been applied to study the effects of cryosurgery in carcinoma of the prostate [52]. Freezing of tissues results in a substantial reduction in T₂*.

Quantitative approaches

Quantitation may include specific MR properties, including, in particular, T₂ and T₂* [53,54], the properties of the remaining signal after long T₂ components are suppressed and the ratio of short T₂ to long T₂ components. There are also other features, such as the magic-angle effect and dipolar contrast, which can be characterized.

There are issues with measuring T₂ and T₂* in the correct range, characterizing different T₂ components (e.g., long and short) including their relative proportions and dealing with artifacts from various sources. Quantitation may be confounded by slice selection and eddy current problems, and by contamination of short T₂ components with long T₂ components that are present in higher concentration.

Future perspective

Imaging of short T₂ and T₂* components is a rapidly expanding area that has seen a convergence of methods targeted at tissues with short T₂ components, susceptibility effects, MIOP imaging and metal artifact control. The methods have improved on solid state imaging, spectroscopy (including continuous-wave methods), electron spin resonance and MR microscopy. The much lower technical performance of clinical systems compared with small-bore spectrometers is a major limitation, but innovative methods for overcoming this problem are now being developed.

The tissues of interest have mainly been in the musculoskeletal system but all tissues of the body have some short T₂ components and study of these may prove to be of diagnostic interest. Some techniques, such as imaging in the presence of metal, are likely to be useful in the clinical domain immediately, while others may require validation and comparative assessment to establish their role. Sodium and phosphorous studies may also be of interest [55,56]. Quantitative approaches may be particularly useful given the large fractional changes in T₂ and T₂* that are frequently seen in disease. The techniques used for imaging often require high gradient performance with control of short-term eddy currents to a level not previously necessary in clinical MR systems. Despite these and other technical difficulties, application to the study of short T₂ and T₂* tissues appears likely to be an area of MRI of considerable importance in the near future.

Financial & competing interests disclosure

The author has received support from General Electric Healthcare. The author has no other relevant affiliations or financial involvement with any organization or entity with a financial interest in or financial conflict with the subject matter or materials discussed in the manuscript apart from those disclosed.

No writing assistance was utilized in the production of this manuscript.

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• Distinction of red and white zones.

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