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Techniques for diffusion-weighted imaging of bone marrow
European Journal of Radiology 55 (2005) 64–73
Techniques for diffusion-weighted imaging of bone marrow J.G. Raya ∗ , O. Dietrich, M.F. Reiser, A. Baur-Melnyk Department of Clinical Radiology, Großhadern, Ludwig Maximilian University, Marchioninistr. 15, 81377 Munich, Germany Received 17 January 2005; received in revised form 20 January 2005; accepted 26 January 2005
Abstract Diffusion-weighted magnetic resonance imaging (DWI) is an imaging technique which is sensitive to random water movements in spatial scales far below those typically accessible by magnetic resonance imaging (MRI). This property makes DWI a powerful tool for diagnosis of diseases which involve alterations in water mobility, such as acute stroke. In bone marrow, DWI has been proven to be a highly useful method for the differential diagnosis of benign and malignant compression fractures. Unfortunately, the application of DWI sequences to the bone marrow frequently suffers from artifacts, which in some cases seriously restrict the diagnostic utility of the image. This requires the introduction of additional correction techniques, or even the development of new sequences. Thus, the selection of an adequate imaging technique for DWI of the bone marrow is a very important issue. In this article the most important sequences for DWI of the bone marrow are reviewed. Special attention is paid to the problems associated with these sequences, as well as their possible solutions. © 2005 Elsevier Ireland Ltd. All rights reserved. Keywords: MRI; Diffusion-weighted imaging; Bone marrow
1. Introduction In the last decades magnetic resonance imaging (MRI) has become an indispensable tool for clinical diagnosis. A major role in its success has been played by its highly non-invasive character and its unique capability to differentiate soft tissues. Tissues can be distinguished in MRI on the basis of their different structural and chemical properties, which determine their behavior in the presence of a magnetic field. For example, in MRI different tissues can be differentiated from each other due to their different magnetization relaxation times. Other contrast mechanisms in MRI are also possible as well. Whenever tissue-dependent changes in the signal intensity are induced, a contrast mechanism can be established. Of course, the usefulness of a particular mechanism depends on its reliability and capability to diagnose diseases whose detection or diagnosis are difficult with the standard techniques. One method which has been shown to be of an exceptional diagnostic value is diffusion-weighted imaging (DWI). This technique is based on the tissue-depending sig∗
Corresponding author. E-mail address: [email protected] (J.G. Raya).
0720-048X/$ – see front matter © 2005 Elsevier Ireland Ltd. All rights reserved. doi:10.1016/j.ejrad.2005.01.014
nal attenuation caused by incoherent thermal motion of water molecules (Brownian motion). Changes of the motion of water molecules in biological tissue can be detected in various pathologic conditions. Initially, DWI was successfully applied in disorders of the brain. The first confirmation of its diagnostic possibilities was in the assessment of acute stroke, which can be detected by DWI within minutes after its onset [1,2]. More recently, DWI has revealed great potential in diagnosing of bone marrow diseases. This technique has been proven to be a highly useful method for the differential diagnosis of acute osteoporotic and neoplastic vertebral compression fractures [3,4]. DWI can distinguish between both types of fractures since the microscopic structure is altered in both cases in a different way. DWI of benign fractures shows iso- or hypointensity in comparison to normal bone marrow because of a higher water mobility in bone marrow edema. Conversely, compression fractures due to malignant tumors lead to hyperintensity related to the high cellularity of tumor tissue. Other applications of DWI of the bone marrow include the detection of infectious diseases [5], and the monitoring of therapy in patients with metastatic infiltration [6].
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Unfortunately, the application of DWI to bone marrow is greatly prone to artifacts. For example, spin echo sequences are normally affected by motion artifacts mostly caused by physiological motions of the patient, such as cardiac and respiratory motion and the pulsation of the cerebrospinal fluid (CSF). These movements can result in severe ghosting artifacts, so that the diagnostic value is no longer given in most of the cases. Therefore, in order to preserve the image quality extra correction techniques or even the development of new sequences are required. One way to eliminate motion artifacts is to use a single-shot sequence, whose very short acquisition time normally underlies the time scale of human physiological motions. However, these sequences are affected by other artifacts, which can also limit their applicability to DWI, and sometimes make the redesign the sequences inevitable. In this article the basic principles of DWI are summarized, and the most common sequences for DWI of the bone marrow are reviewed in detail. Special attention is paid to the explanation of the most common artifacts of each sequence type and to possible solutions how these artifacts might be overcome.
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Due to the random displacements of water during the application of magnetic field gradients each molecule acquires a different, path-dependent phase. In practice, this phase spreading causes a loss of spin precession coherence at the time of signal acquisition and, therefore, a decrease of the signal intensity. The influence of self-diffusion on spin echo amplitudes was already described by Hahn in his 1950 original paper [7]. More than 30 years later, the diffusion effects were demonstrated to offer great potentials in clinical MRI [1,8,9]. Since the ADC strongly depends on the tissue properties, it is possible to establish a contrast mechanism based on ADC differences. Such techniques for diffusion-contrasted imaging have become broadly known as diffusion-weighted imaging (DWI). To profit from the tissue-depending signal attenuation, it is useful to feature a method for the determination of the ADC. The method almost all diffusion sequences are based on was proposed in 1965 by Stejskal and Tanner [10] as pulsed-gradient spin echo (PGSE) measurement. The idea was to use two identical gradients, the so-called diffusion gradients, at both sides of a 180◦ refocusing pulse, in order to separate the diffusion-sensitizing from other sequence parameters. In this case the signal shows an exponential decay,
2. Physical principles of DWI MRI in vivo is usually based on the coherent precession of the proton spins of water in an external magnetic field due to the abundance of hydrogen atoms in living organisms. An important point concerning the formation of the MRI signal is that by the presence of magnetic field gradients random movements of molecules containing hydrogen produce a partial annihilation of the precession coherence of the spins (so-called de-phasing), and thus a decrease of the signal intensity results. Thermal fluctuations are responsible for keeping water molecules in random motion. This thermal motion is technically known as (self-)diffusion or as Brownian molecular motion, and is a typical stochastic phenomenon, which can be well described with the help of statistical mechanics. Furthermore, since in clinical routine MRI offers a maximum spatial resolution of the order of a few hundred micrometers, the detailed movement of a single water molecule cannot be resolved. Therefore, only averaged diffusion quantities are relevant for MRI. The most important physical quantity is the diffusion coefficient (or diffusivity), D, which is proportional to the mean-squared value of the displacement per unit time (mm2 /s). The actual value of the diffusivity strongly depends on the environment or media in which the water molecules move. In particular, the structure of the different biological tissues sets some characteristic natural barriers to the free spread of water. These tissue dependent constrains are reflected in the value of the diffusion coefficient, which becomes different from that of free water. Therefore the diffusivity in tissue is usually called apparent diffusion coefficient (ADC).
I = I0 exp(−b ADC),
(1)
where I is the intensity in the presence of the diffusion gradients, I0 the intensity in the absence of diffusion gradients and b the so-called b-value, b = (γgδ)
2
δ − , 3
(2)
which only depends on the gyromagnetic ratio γ, and on diffusion gradient parameters: their amplitude, g, their duration, δ, and the separation time between the onset of the two diffusion gradients, · This exponential dependence allows an easy determination of the ADC. Chronologically, the spin echo (SE) was the first NMR pulse sequence used to measure diffusivity [7]. Remarkably, this happened many years before the technique of MRI was developed. In DWI the ADC is determined as follows. First of all, several (at least two) images of the same slice are taken with different b-values. Then the ADC is calculated by performing an exponential fit or a linear regression analysis either to the mean intensity of a region of interest or pixelby-pixel to the whole image. In the latter case, the resulting image, in which each pixel represents an ADC value, is the so-called ADC map. Determination of the apparent diffusivity using the wellestablished techniques of MRI has been shown to be a useful diagnostic tool. Nowadays, it has become a standard procedure in clinical practice for the diagnosis of many diseases.
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3. Diffusion-weighted sequences for bone marrow imaging The great versatility and simplicity of the Stejskal-Tanner method made it possible to adapt numerous MRI sequences to DWI. However, the addition of extra diffusion gradients is not straightforward for all sequence types, and therefore is usually implemented for each sequence type in a different way. Furthermore, the presence of such gradients is normally associated with some undesirable effects that seriously degrade the image quality. The most common are the so-called motion artifacts, which result from patient movements during the diffusion-sensitizing interval. Motion artifacts are specially problematic in imaging of bone marrow, since the anatomical structures containing bone marrow (such as the spine) are normally prone to physiological motions like the cerebrofluid pulsation. As a consequence of the difficulties mentioned above in adapting MRI sequences to DWI a detailed discussion for the different sequences is necessary. First of all, a specific problem appears due to the composition of bone marrow, which is relatively poor in water and rich in fat. Thus DWI will show a water signal with the intensity depending on the b-value superimposed by a b-independent high signal coming from fat. Therefore, it is recommendable to suppress the fat signal by applying one of the established
fat suppression techniques such as RF fat saturation or shorttau inversion recovery (STIR) imaging. In the following sequences which are commonly used in DWI of the bone marrow are described. First of all SE-based sequences are reviewed. Common to most of these sequences is a relatively high sensitivity to patient movements, which normally requires additional technical developments in order to correct the image from artifacts. Fast acquisition sequences allow to obtain images almost without any motion artifacts. However, they are prone to other artifacts, which can also seriously degrade the quality of the image. Among the fast acquisition techniques we will focus on the turbo spin echo (TSE)-based sequences, on diffusion-weighted echo planar imaging (EPI), and on steady-state free precession (SSFP) sequences. 3.1. Spin echo-based sequences for diffusion-weighted imaging Both the spin echo (SE) and the stimulated echo (STE) sequences have played a pioneering role in DWI, since for both techniques the Stejskal-Tanner method can be easily applied (see Fig. 1A). Although these sequences offer some advantages, they also have serious disadvantages. Among their positive features are their relatively high signal-to-noise ratio (SNR) and their low sensitivity with respect to inhomo-
Fig. 1. Sequence scheme of different SE-based sequences. In each scheme only a single pulse sequence excitation cycle is represented. The diffusion gradients (in these examples applied in readout direction) are always drawn with a diagonal hatch. (A) Diffusion-weighted spin echo sequence. (B) Navigator-corrected spin echo diffusion sequence. The second refocusing pulse produces a second echo, the navigator echo, which is not phase-encoded. (C) Radial k-space spin echo diffusion sequence. Varying the direction of the readout gradients, different radial trajectories in k-spaces are acquired. (D) Line scan diffusion sequence. The slice selection gradients of the excitation and refocusing pulse are inversely rotated around the readout direction in order to produce an echo of only a single (one-dimensional) line.
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geneities in the susceptibility of the measured object. However, these sequences require long acquisition times. Because of the long longitudinal relaxation time of most tissues, it is necessary to wait long enough before the next exciting pulse can be applied. Otherwise, the signal will be very low in the following acquisition. However, the most limiting characteristic is their high sensitivity to coherent macroscopic motions. As a consequence of the large duration and amplitude of the diffusion gradients all moving protons acquire an identical additional extra phase during each repetition time. Due to the properties of the Fourier transform, this extra phase does
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not lead to a higher attenuation of the signal, but to a spatial translation of the reconstructed image. During each repetition a different global phase is added and signal from the entire slice is received. Thus the image appears as a superposition of multiple similar images each one displaced parallel to the other by a different amount. The resulting ghosting artifacts which spread in phase-encoding direction make the image useless for diagnostic purposes (Fig. 2A). Different methods have been proposed to overcome the limitations which are imposed on the SE and STE sequences by motion artifacts. One of these is the navigator echo tech-
Fig. 2. An example of motion artifacts and the way of correction. (A) An image obtained from a volunteer with a diffusion-weighted SE sequence. Motion artifacts make this image non-diagnostic. (B) Motion artifacts are partially corrected using extra gating. Thus a reduced ghosting in the area of the cerebrospinal fluid can be observed. (C) The use of navigator echoes gives rise to a considerable reduction of artifacts in the upper part of the image. However, since this image is not triggered, artifacts at the height of the lateral ventricles persist. (D) A combination of both correction mechanisms result in an optimal image quality. Interestingly, a residual motion artifact caused by the eyes motions (see white arrow) can be also observed in image.
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nique, which was independently proposed by Ordidge et al. [11] and Anderson and Gore [12] in 1994 (see Fig. 1B). The basic idea is to measure and correct the motion-induced extra phases. With this aim an additional echo, known as navigator echo, is acquired without phase-encoding after each acquisition. In the absence of motion all the navigator echoes should be identical. Motion will affect in the same way both the phase-encoded echo and the navigator echo. Therefore, the differences between two successive navigator echoes allows one to calculate the relative extra phase added by the motion. By the subtraction of these phases most of the motion artifacts are suppressed (see Fig. 2C). Unfortunately, not all motion can be corrected to the same extent. For example, in DWI of the spine, cerebrospinal fluid pulsations still lead to artifacts after phase correction (see Fig. 2B). Therefore, an additional gating is usually needed in order to synchronize acquisitions with pulsations (see Fig. 2D). The combination of all these correction techniques allows to obtain almost artifact free images of the bone marrow [13]. However, such a procedure presents two important disadvantages. On the one hand, it is a technically complex method, and on the other hand, the acquisition time is considerably long. All these facts complicate the integration in the clinical routine. In order to increase the robustness of navigated diffusionweighted SE sequences, radial k-space SE sequences have been developed [14] (see Fig. 1C). This sequence type offers the advantage of self-navigation, since the center of the k-space is read in each line. Therefore, no additional echo for the phase correction is needed. Moreover, a selection algorithm, which discards the lines affected by complex motions such as regional and pulsating motions, reduces artifacts. The residual effects of cardiac motion and CSF pulsations also have to be suppressed by means of cardiac gating. This sequence has been demonstrated as a robust technique for high quality imaging of the spine [14]. Gudbjartsson et al. [15] have adopted a different approach in their line scan diffusion technique, which is a variant of the SE sequence (see Fig. 1D). They proposed to construct the image line by line from one-dimensionally acquired dataset, so that the relative phase between adjacent lines does not play any role. This sequence is inherently insensitive to motion and to susceptibility artifacts. In addition, due to a special selection scheme of the lines it is no longer necessary to wait for longitudinal relaxation before the next excitation, making this sequence considerably faster than the SE sequence. However, since the echo is produced only in a line and not in the whole slice, this sequence has a low SNR and therefore requires a careful optimization in order to keep an acceptable signal level. Another possibility of overcoming motion effects is to drastically decrease the acquisition time. MRI offers various fast acquisition techniques which can also be applied in DWI. Many of them are based on the successive refocusing of a diffusion-weighted echo, so that a complete image can be read after only one excitation. The intrinsic
transversal relaxation times, T2 or T2∗ , impose an unavoidable constraint to the time in which an echo can be refocused and, therefore, to the achievable spatial resolution of these sequences. 3.2. TSE-based diffusion weighted-sequences One way of refocusing the spins is by means of 180◦ pulses. The most relevant example of such a sequence is the turbo spin echo (TSE) sequence with the single-shot variants “rapid acquisition with relaxation enhancement” (RARE) [16] and “half-Fourier acquired single-shot turbo spin echo” (HASTE) [17]. These sequences have been adapted to DWI in spite of their sensitivity to the magnetization phase prior to the train of refocusing pulses, i.e. after its preparation. Small phase shifts of the magnetization with respect to the inversion axis of the 180◦ pulse lead to image artifacts. Therefore, a diffusion-weighted preparation of the magnetization, which inherently involves phase shift and spread, is expected to cause severe artifacts. However, appropriate modifications of these sequences allow to make them insensitive to the initial phase of the magnetization so that they become suitable for DWI. In order to modify the TSE-based sequences for DWI, a good understanding of the echo formation mechanism is needed. A central role plays the non-ideal character of the 180◦ refocusing pulse, which rotate the spins across the slice profile not exactly by 180◦ but by angles from 0 to 180◦ . Those spins which are not exactly inverted lead to stimulated and higher order echoes [18]. Therefore, the spins contributing to an echo, except from the first one, have been de-phased and re-phased in very different ways. Each one of these possible refocusing ways is known as a coherence pathway, and can be characterized according to the parity of the number of inversions that the spins have undergone. Spins following an even-parity pathway undergo an even number of inversions of the magnetization around the pulse direction, and thus the magnetization at the echo time keeps its initial orientation. However, spins in an odd-parity pathway suffer an odd number of inversions, thus leading to an inverted magnetization with respect to the inversion axis of the pulse. Since all spins contributing to an echo refocus at the same time, the component of the transversal magnetization which does not lie parallel to the axis of the refocusing pulse is progressively annihilated as the echo number increases and both types of coherence pathways turn to be equally populated. Two possibilities exist to avoid the destructive addition of coherent pathways. The first one is to split the even- and odd-parity echoes. A second solution is the elimination of the non-parallel component of the transversal magnetization prior to the first refocusing pulse. The “displaced ultra-fast low angle RARE sequence” (displaced U-FLARE) [19] belongs to the first group. In this sequence an extra de-phasing gradient is added before the readout gradient takes place (see Fig. 3A). Spins, which contribute to an echo following an even-parity pathway, add no net phase due to the new gradi-
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Fig. 3. Sequence schema of displaced U-FLARE, EPI and PSIF sequences. The diffusion gradients are always drawn with a diagonal hatch. (A) Displaced U-FLARE sequence. The part of the sequence outside the brackets prepares the transversal magnetization. The sequence part between brackets is responsible for the echo-refocusing, and is repeated until all the lines are acquired. The gradient in black is the de-phasing gradient introduced to split even- and odd-parity echoes. (B) Scheme of an echo-planar imaging sequence. As in A, the part of the sequence outside the brackets prepares the diffusion-weighting of the magnetization. Between the brackets is the part of the sequence that refocuses the echo. Each of the gradients in the readout direction allows the acquisition of an echo. After each echo a phase-encoding is performed by the small gradient in black (the so-called “blip”). (C) PSIF sequence. An excitation of a small angle, θ, is continuously repeated in order to achieve a steady-state of the magnetization. In each excitation a diffusion gradient is applied.
ent. On the contrary, all the spins in an odd-parity pathway acquire an extra phase. Thus echoes split in time according to their parity. Strong enough de-phasing gradients allow a shift of odd-echoes out of the acquisition window. A disadvantage of this procedure is presented by the signal intensity being reduced by a factor of two, since only the half of the echoes are acquired (see Fig. 4B for an example). Moreover, the maximum signal intensity is achieved after a few refocusing pulses, which implies that the first echoes must be discarded. Schick [18] proposed a sequence, “the split acquisition of fast spin echo signals for diffusion imaging” (SPLICE), which solve the SNR problem. In this sequence an unbalanced readout gradient is responsible for the splitting of the echoes, which are acquired separately. A final addition of the Fourier-transformed amplitude images of each echo results in an improvement of the SNR. A disadvantage of the SPLICE sequence is the increased echo spacing and thus the prolonged echo train duration.
Among the sequences in which the non-parallel component of the magnetization is eliminated, the “phaseinsensitive RARE pulse sequence” [20] introduced by Alsop stands out. Inserting a 90◦ pulse at one half echo time before the first refocusing pulse removes the out of phase component of the transversal magnetization which avoids the problem of destructive superposition of coherence pathways. Even more, the use of a tailored train of RF pulses produces a stable echo amplitude from the first echo and drastically diminishes the power deposition. The main disadvantage of this technique is that the signal is reduced by a factor of two due to the phase insensitive preparation of the magnetization. 3.3. Diffusion-weighted echo-planar imaging Another way of refocusing the spins is to use gradient echoes instead of spin echoes. The most relevant single-shot sequence by using a train of pure gradient echoes is the so-
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Fig. 4. Comparison of EPI, RARE and PSIF images obtained in the same patient. (A) 128 × 96 image acquired with an EPI sequence. As can be appreciated here, EPI images are normally degraded by artifacts, which seriously limit the clinical usefulness of the image. Some of the artifacts are caused by chemical shift (in this example, the structure indicated with arrow 1 may be a chemical shift artifact), and susceptibility inhomogeneities (see the distortion indicated with arrow 2). (B) 128 × 96 image acquired with a RARE sequence. Image quality is suitable for clinical purposes, though the image suffers from blurring. (C) 512 × 512 image interpolated from a 256 × 256 image acquired with a PSIF sequence. This high resolution image shows a sharp contrast and almost no artifacts. Thus PSIF images are very appropriate for clinical applications. The disadvantage of this sequence is that no quantification of the ADC can be performed.
called echo-planar imaging (EPI) sequence [21]. The adaptation of the spin echo EPI sequences to DWI requires a pair of diffusion gradients before the first acquisition (see Fig. 3B). Unlike the TSE sequence the extra addition of the diffusion gradients does not produce any undesired echo modulation and image artifacts. However, DW-EPI causes the same image artifacts as EPI [22], thus seriously restricting the applicability of DW-EPI for bone marrow imaging (see Fig. 4A). The most limiting artifacts are caused by off-resonance effects such as susceptibility artifacts. Common to all off-resonance effects is that not all the spins possess the same resonance (Larmor) frequency. To understand the implications of a resonance frequency spread for an EPI measurement let us consider an ideal experiment (see Fig. 5). Two spin species, A and B, with slightly different resonance frequencies are isotropically distributed in a cylindrical phantom, which lies parallel to a perfectly homogeneous external magnetic field. An ideal EPI experiment is performed on this phantom. First a 90◦ RF pulse turns over all the spins, which are included in a transversal slice located at one half of the phantom height. The flipped atoms of both species, initially in phase, start precessing. Since in the gradient echo technique the echo is produced without inverting the spins, differences in the resonance frequency produce a continuously increasing relative phase between both species, which otherwise re-phase at the same time (see in Fig. 5 the last two boxes). As it has been already seen, an additional phase will result in displacements in the reconstructed image. Therefore, the images produced from the species A and B will be displaced from each other (see the idealized imaging in Fig. 5). The actual way in which both images are displaced depends on
the accumulated phase out, i.e. on the way the k-space is sampled. In an EPI measurement the k-space is sampled in such a way that the points in readout direction are acquired very rapidly, whereas the sampling in the phase-encoding direction takes place more slowly, i.e. with a much lower effective receiver bandwidth (e.g. for a 128 × 128 acquisition matrix the sampling rate in phase-encoding direction is approximately by a factor of 128 greater than in the read out direction). Therefore, the accumulated phase difference between the two species is larger between two points in the phase-encoding direction than in the readout direction. Even more, since the phase difference grows at an uniform rate, the phase encoding is not altered and thus no smearing appears in the image. Obviously, the displacement caused by the accumulated phase difference must be compared to the pixel size in order to see whether visible artifacts occur. Normally, owing to the different acquisition rates in phase-encoding and readout directions, no artifacts can be seen in readout direction whereas, even by very small off-resonance effects (∼3 ppm), displacements in the phase-encoding direction are in the order of several pixels. Off-resonance effects are caused by distortions in the external magnetic field, chemical shift, or susceptibility inhomogeneities. With an adequate shimming it is possible to homogenize the external magnetic field in order to eliminate any artifact. Chemical shift can be reduced using a fat suppression or water-only excitation. This fact is of special importance in imaging of bone marrow due to the high content of fat within the yellow marrow. Imperfect fat suppression would lead to residual artifacts. The susceptibility in-
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homogeneities are the most problematic off-resonance effect when performing EPI in vivo, since a correction can hardly be performed. Susceptibility inhomogeneities mostly take place in the vicinity of interfaces between different media such as in the interfaces between bone, soft tissue, and air around the spine. Therefore, the applicability of DWEPI sequences to the bone marrow is seriously limited (see Fig. 4A). 3.4. Multi-shot echo-planar imaging A simple way to overcome the limitations that the T2∗ relaxation time imposes on the maximal echo train length and thereby, on the spatial resolution of EPI images is just to cover the complete k-space with different excitations and to combine all the data in a single raw data set [22]. Since each excitation requires a diffusion-weighting, patient motions can cause phase errors between the different data sets. Therefore, a navigated phase-correction is in the majority of cases necessary in order to prevent from motion artifacts. This technique has been proved to produce images of high diagnostic quality [22]. One of the major problems in breaking up the k-space is that phase and amplitude discontinuities between different data sets produce blurring and ghosting. Therefore, the most robust method to read out the k-space with a multi-shot EPI sequence is to acquire the k-space lines interleaved. With an adequate selection of the RF pulse flip angles a smooth variation of the signal intensities and phase throughout the k-space can be achieved avoiding artifacts [22]. Even more, this method is less sensitive to off-resonance effects. Since successively acquired lines are now separated by a number of lines equal to the number of excitations, Ne , the accumulated phase difference between spins in phase-encoding direction
Fig. 5. Ideal measurement with an EPI sequence in order to illustrate offresonance effects. A cylindrical phantom filled up with two spin species, A and B, with slightly different resonance frequencies lies parallel to a perfect homogeneous external magnetic field. An ideal EPI measurement, as depicted in the figure, is perform. In the boxes the behavior of the spins of type A and B in the rotating reference system (at the frequency of atoms of type A) is shown at different moments of the measurement. The spins are supposed to lie in the readout direction. Relaxation effects have not been considered in this idealized experiment. In the first box from the left the spins are represented just after the application of the 90◦ RF pulse. Both species are initially in phase. In the second box spins are represented while the de-phasing gradient is switched on. Because of the gradient in the readout direction, spins located at different positions on the readout direction accumulate a different phase. Due to their different resonance frequencies spins of type A and B have begun to de-phase (the white sector represents the phase difference between A and B, which is the same for all spins). Since the gradient echo does not turn over the spins, the phase difference continues growing through the measurement. Boxes three and four show how all the spins re-phase at the same time but with a growing phase difference (white sector). At the bottom the ideally resulting image is represented. The accumulated phase difference of B-spins causes a displacement of their image in the phase direction.
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is divided by Ne . Therefore, off-resonance effects in phase direction are reduced, thus making this sequence much more appropriate for DWI of the bone marrow than DW-EPI. 3.5. Steady-state free precession sequences A different type of fast acquisition sequences are those based on creating a steady-state of the longitudinal and transversal magnetization, the so-called steady-state free precession (SSFP) sequences. In these sequences the magnetization is repeatedly disturbed by a fast repetition of RF pulses. After some RF pulses the transversal magnetization
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achieves a steady-state, in which the amount of longitudinal and transversal magnetization after each pulse is the same. SSFP sequences can be adapted to DWI by introducing a diffusion gradient in each inter-pulse period (see Fig. 3C). The most useful of these sequences for the bone marrow is the diffusion-weighted PSIF sequence [23] (reversed fast imaging with steady precession (FISP [24]) sequence), which allows the acquisition of high resolution images in times of the order of few seconds because of the short repetition times of less than 50 ms. Since the RF pulse flip angle are always smaller than 180◦ , each pulse turns a part of the longitudinal magnetization into the transversal plane and partially brings the transversal magnetization back parallel to the magnetic field. Therefore, as in the TSE-based sequences, the echo is formed by the superposition of multiple coherence pathways. However, unless in the TSE-based sequences, every pathway has a different diffusion-weighting. The signal intensity depends on the sequence parameters and the diffusivity in such a very complicated way, that it is not suitable for quantitative determination of any physical constant [25]. Nevertheless, the good quality of the images showing almost no artifacts makes this sequence a very useful tool for diagnostics (see Fig. 4C).
4. Conclusions In this article we describe the different sequences that have been used for DWI of the bone marrow so far. The problems of each sequence type is analyzed carefully and their possible solutions are explained. Although the standard diffusionweighted SE and TSE pulse sequences are seriously affected by motion artifacts, images suitable for clinical use can be obtained either by adding a correction technique such as the navigator echo, or by using new developed SE-based sequence (such as the k-space spin echo or the line scan diffusion imaging sequences). Due to the bone marrow composition and location, single-shot sequences result in images which do not always offer the desired quality. For example, susceptibility artifacts of DW-EPI sequences seriously restrict the image quality. This problem is partially solved by the multi-shot DW-EPI sequence. SSFP sequences allow the acquisition of good quality images in short time, but with the disadvantage of not being appropriate for quantitative DWI. Unlike in the imaging of the brain, where DW-EPI has been established as the standard method for DWI, no generally accepted standard sequence for DWI of the bone marrow has been established.
Acknowledgment The authors would like to thank Dr. Sonja Buhmann for providing the MR images of Fig. 4 and for critical reading of the manuscript. This work was supported by
the Deutsche Forschungsgemeinschaft (DFG), Grant No. BA-2089(1-3).
References [1] Le Bihan D, Breton E, Lallemand D, et al. MR imaging of intravoxel incoherent motions: application to diffusion and perfusion in neurologic disorders. Radiology 1986;161:401–7. [2] Mintorovitch J, Moseley ME, Chileuitt L, et al. Comparison of diffusion- and T2-weighted MRI for the early detection of cerebral ischemia and reperfusion in rats. Magn Reson Med 1991;18: 39–50. [3] Baur A, St¨abler A, Br¨uning R, et al. Diffusion-weighted MR imaging of bone marrow: differentiation of benign versus pathologic compression fractures. Radiology 1998;207:349–56. [4] Baur A, Huber A, Ertl-Wagner B, et al. Diagnostic value of increased diffusion weighting of a steady-state free precession sequence for differentiating acute benign osteoporotic fractures from pathologic vertebral compression fractures. Am J Neuroradiol 2001;22: 366–72. [5] Byun WM. Diffusion-weighted MR imaging of vertebral bone marrow: differentiation of degenerative spines and spondylitis involving bone marrow adjacents to end plates. In: Books of Abstracts. Glasgow: International Society of Magnetic Resonance in Medicine; 2001. p. 1626. [6] Byun WM, Shin SO, Chang Y, et al. Diffusion-weighted MR imaging of metastatic disease of the spine: assessment of response to therapy. Am J Neuroradiol 2002;23:906–12. [7] Hahn EL. Spin echoes. Phys Rev 1950;80:580–94. [8] Merboldt KD, H¨anicke W, Frahm J. Self-diffusion NMR imaging using stimulated echoes. J Magn Reson 1985;64:479–86. [9] Taylor DG, Bushell MC. The spatial mapping of translational diffusion coefficients by the NMR imaging technique. Phys Med Biol 1985;30:345–9. [10] Stejskal EO, Tanner JE. Spin diffusion measurements: spin echoes in the presence of time-dependent field gradient. J Chem Phys 1965;42:288–92. [11] Ordidge RJ, Helpern JA, Qing ZX, et al. Correction of motional artifacts in diffusion-weighted MR images using navigator echoes. Magn Reson Imag 1994;12:455–60. [12] Anderson AW, Gore JC. Analysis and correction of motion artifacts in diffusion weighted imaging. Magn Reson Med 1994;32:379– 87. [13] Dietrich O, Heiland S, Benner T, et al. Reducing motion artifacts in diffusion-weighted MRI of the brain: efficacy of navigator echo correction and pulse triggering. Neuroradiology 2000;42:85–91. [14] Dietrich O, Herlihy A, Dannels WR, et al. Diffusion-weighted imaging of the spine using radial k-space trajectories. MAGMA 2001;12:23–31. [15] Gudbjartsson H, Maier SE, Mulkern RV, et al. Line scan diffusion imaging. Magn Reson Med 1996;36:509–11. [16] Henning J, Nauerth A, Friedburg H. RARE imaging: a fast imaging method for clinical MR. Magn Reson Med 1986;3:823–33. [17] Kiefer B, Gr¨assner J, Hausmann R. Image acquisition in a second with half-Fourier acquired single shot turbo spin echo. Magn Reson Imag 1994;4:86–92. [18] Schick F. SPLICE: sub-second diffusion-sensitive MR imaging using a modified fast spin-echo acquisition mode. Magn Reson Med 1997;38:638–44. [19] Norris DG. Ultrafast low-angle RARE: U-FLARE. Magn Reson Med 1991;17:539–42. [20] Alsop DC. Phase insensitive preparation of single-shot RARE: application to diffusion imaging in humans. Magn Reson Med 1997;38:527–33.
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[24] Oppelt A, Graumann R, Borfuss H, et al. FISP: a new fast MRI sequence. Electromedica 1986;54:15–8. [25] Deimling M, Heid O. High resolution SSFP diffusion imaging. In: Book of Abstracts. San Francisco: International Society of Magnetic Resonance in Medicine; 1994. p. 1033.
Techniques for diffusion-weighted imaging of bone marrow J.G. Raya ∗ , O. Dietrich, M.F. Reiser, A. Baur-Melnyk Department of Clinical Radiology, Großhadern, Ludwig Maximilian University, Marchioninistr. 15, 81377 Munich, Germany Received 17 January 2005; received in revised form 20 January 2005; accepted 26 January 2005
Abstract Diffusion-weighted magnetic resonance imaging (DWI) is an imaging technique which is sensitive to random water movements in spatial scales far below those typically accessible by magnetic resonance imaging (MRI). This property makes DWI a powerful tool for diagnosis of diseases which involve alterations in water mobility, such as acute stroke. In bone marrow, DWI has been proven to be a highly useful method for the differential diagnosis of benign and malignant compression fractures. Unfortunately, the application of DWI sequences to the bone marrow frequently suffers from artifacts, which in some cases seriously restrict the diagnostic utility of the image. This requires the introduction of additional correction techniques, or even the development of new sequences. Thus, the selection of an adequate imaging technique for DWI of the bone marrow is a very important issue. In this article the most important sequences for DWI of the bone marrow are reviewed. Special attention is paid to the problems associated with these sequences, as well as their possible solutions. © 2005 Elsevier Ireland Ltd. All rights reserved. Keywords: MRI; Diffusion-weighted imaging; Bone marrow
1. Introduction In the last decades magnetic resonance imaging (MRI) has become an indispensable tool for clinical diagnosis. A major role in its success has been played by its highly non-invasive character and its unique capability to differentiate soft tissues. Tissues can be distinguished in MRI on the basis of their different structural and chemical properties, which determine their behavior in the presence of a magnetic field. For example, in MRI different tissues can be differentiated from each other due to their different magnetization relaxation times. Other contrast mechanisms in MRI are also possible as well. Whenever tissue-dependent changes in the signal intensity are induced, a contrast mechanism can be established. Of course, the usefulness of a particular mechanism depends on its reliability and capability to diagnose diseases whose detection or diagnosis are difficult with the standard techniques. One method which has been shown to be of an exceptional diagnostic value is diffusion-weighted imaging (DWI). This technique is based on the tissue-depending sig∗
Corresponding author. E-mail address: [email protected] (J.G. Raya).
0720-048X/$ – see front matter © 2005 Elsevier Ireland Ltd. All rights reserved. doi:10.1016/j.ejrad.2005.01.014
nal attenuation caused by incoherent thermal motion of water molecules (Brownian motion). Changes of the motion of water molecules in biological tissue can be detected in various pathologic conditions. Initially, DWI was successfully applied in disorders of the brain. The first confirmation of its diagnostic possibilities was in the assessment of acute stroke, which can be detected by DWI within minutes after its onset [1,2]. More recently, DWI has revealed great potential in diagnosing of bone marrow diseases. This technique has been proven to be a highly useful method for the differential diagnosis of acute osteoporotic and neoplastic vertebral compression fractures [3,4]. DWI can distinguish between both types of fractures since the microscopic structure is altered in both cases in a different way. DWI of benign fractures shows iso- or hypointensity in comparison to normal bone marrow because of a higher water mobility in bone marrow edema. Conversely, compression fractures due to malignant tumors lead to hyperintensity related to the high cellularity of tumor tissue. Other applications of DWI of the bone marrow include the detection of infectious diseases [5], and the monitoring of therapy in patients with metastatic infiltration [6].
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Unfortunately, the application of DWI to bone marrow is greatly prone to artifacts. For example, spin echo sequences are normally affected by motion artifacts mostly caused by physiological motions of the patient, such as cardiac and respiratory motion and the pulsation of the cerebrospinal fluid (CSF). These movements can result in severe ghosting artifacts, so that the diagnostic value is no longer given in most of the cases. Therefore, in order to preserve the image quality extra correction techniques or even the development of new sequences are required. One way to eliminate motion artifacts is to use a single-shot sequence, whose very short acquisition time normally underlies the time scale of human physiological motions. However, these sequences are affected by other artifacts, which can also limit their applicability to DWI, and sometimes make the redesign the sequences inevitable. In this article the basic principles of DWI are summarized, and the most common sequences for DWI of the bone marrow are reviewed in detail. Special attention is paid to the explanation of the most common artifacts of each sequence type and to possible solutions how these artifacts might be overcome.
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Due to the random displacements of water during the application of magnetic field gradients each molecule acquires a different, path-dependent phase. In practice, this phase spreading causes a loss of spin precession coherence at the time of signal acquisition and, therefore, a decrease of the signal intensity. The influence of self-diffusion on spin echo amplitudes was already described by Hahn in his 1950 original paper [7]. More than 30 years later, the diffusion effects were demonstrated to offer great potentials in clinical MRI [1,8,9]. Since the ADC strongly depends on the tissue properties, it is possible to establish a contrast mechanism based on ADC differences. Such techniques for diffusion-contrasted imaging have become broadly known as diffusion-weighted imaging (DWI). To profit from the tissue-depending signal attenuation, it is useful to feature a method for the determination of the ADC. The method almost all diffusion sequences are based on was proposed in 1965 by Stejskal and Tanner [10] as pulsed-gradient spin echo (PGSE) measurement. The idea was to use two identical gradients, the so-called diffusion gradients, at both sides of a 180◦ refocusing pulse, in order to separate the diffusion-sensitizing from other sequence parameters. In this case the signal shows an exponential decay,
2. Physical principles of DWI MRI in vivo is usually based on the coherent precession of the proton spins of water in an external magnetic field due to the abundance of hydrogen atoms in living organisms. An important point concerning the formation of the MRI signal is that by the presence of magnetic field gradients random movements of molecules containing hydrogen produce a partial annihilation of the precession coherence of the spins (so-called de-phasing), and thus a decrease of the signal intensity results. Thermal fluctuations are responsible for keeping water molecules in random motion. This thermal motion is technically known as (self-)diffusion or as Brownian molecular motion, and is a typical stochastic phenomenon, which can be well described with the help of statistical mechanics. Furthermore, since in clinical routine MRI offers a maximum spatial resolution of the order of a few hundred micrometers, the detailed movement of a single water molecule cannot be resolved. Therefore, only averaged diffusion quantities are relevant for MRI. The most important physical quantity is the diffusion coefficient (or diffusivity), D, which is proportional to the mean-squared value of the displacement per unit time (mm2 /s). The actual value of the diffusivity strongly depends on the environment or media in which the water molecules move. In particular, the structure of the different biological tissues sets some characteristic natural barriers to the free spread of water. These tissue dependent constrains are reflected in the value of the diffusion coefficient, which becomes different from that of free water. Therefore the diffusivity in tissue is usually called apparent diffusion coefficient (ADC).
I = I0 exp(−b ADC),
(1)
where I is the intensity in the presence of the diffusion gradients, I0 the intensity in the absence of diffusion gradients and b the so-called b-value, b = (γgδ)
2
δ − , 3
(2)
which only depends on the gyromagnetic ratio γ, and on diffusion gradient parameters: their amplitude, g, their duration, δ, and the separation time between the onset of the two diffusion gradients, · This exponential dependence allows an easy determination of the ADC. Chronologically, the spin echo (SE) was the first NMR pulse sequence used to measure diffusivity [7]. Remarkably, this happened many years before the technique of MRI was developed. In DWI the ADC is determined as follows. First of all, several (at least two) images of the same slice are taken with different b-values. Then the ADC is calculated by performing an exponential fit or a linear regression analysis either to the mean intensity of a region of interest or pixelby-pixel to the whole image. In the latter case, the resulting image, in which each pixel represents an ADC value, is the so-called ADC map. Determination of the apparent diffusivity using the wellestablished techniques of MRI has been shown to be a useful diagnostic tool. Nowadays, it has become a standard procedure in clinical practice for the diagnosis of many diseases.
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3. Diffusion-weighted sequences for bone marrow imaging The great versatility and simplicity of the Stejskal-Tanner method made it possible to adapt numerous MRI sequences to DWI. However, the addition of extra diffusion gradients is not straightforward for all sequence types, and therefore is usually implemented for each sequence type in a different way. Furthermore, the presence of such gradients is normally associated with some undesirable effects that seriously degrade the image quality. The most common are the so-called motion artifacts, which result from patient movements during the diffusion-sensitizing interval. Motion artifacts are specially problematic in imaging of bone marrow, since the anatomical structures containing bone marrow (such as the spine) are normally prone to physiological motions like the cerebrofluid pulsation. As a consequence of the difficulties mentioned above in adapting MRI sequences to DWI a detailed discussion for the different sequences is necessary. First of all, a specific problem appears due to the composition of bone marrow, which is relatively poor in water and rich in fat. Thus DWI will show a water signal with the intensity depending on the b-value superimposed by a b-independent high signal coming from fat. Therefore, it is recommendable to suppress the fat signal by applying one of the established
fat suppression techniques such as RF fat saturation or shorttau inversion recovery (STIR) imaging. In the following sequences which are commonly used in DWI of the bone marrow are described. First of all SE-based sequences are reviewed. Common to most of these sequences is a relatively high sensitivity to patient movements, which normally requires additional technical developments in order to correct the image from artifacts. Fast acquisition sequences allow to obtain images almost without any motion artifacts. However, they are prone to other artifacts, which can also seriously degrade the quality of the image. Among the fast acquisition techniques we will focus on the turbo spin echo (TSE)-based sequences, on diffusion-weighted echo planar imaging (EPI), and on steady-state free precession (SSFP) sequences. 3.1. Spin echo-based sequences for diffusion-weighted imaging Both the spin echo (SE) and the stimulated echo (STE) sequences have played a pioneering role in DWI, since for both techniques the Stejskal-Tanner method can be easily applied (see Fig. 1A). Although these sequences offer some advantages, they also have serious disadvantages. Among their positive features are their relatively high signal-to-noise ratio (SNR) and their low sensitivity with respect to inhomo-
Fig. 1. Sequence scheme of different SE-based sequences. In each scheme only a single pulse sequence excitation cycle is represented. The diffusion gradients (in these examples applied in readout direction) are always drawn with a diagonal hatch. (A) Diffusion-weighted spin echo sequence. (B) Navigator-corrected spin echo diffusion sequence. The second refocusing pulse produces a second echo, the navigator echo, which is not phase-encoded. (C) Radial k-space spin echo diffusion sequence. Varying the direction of the readout gradients, different radial trajectories in k-spaces are acquired. (D) Line scan diffusion sequence. The slice selection gradients of the excitation and refocusing pulse are inversely rotated around the readout direction in order to produce an echo of only a single (one-dimensional) line.
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geneities in the susceptibility of the measured object. However, these sequences require long acquisition times. Because of the long longitudinal relaxation time of most tissues, it is necessary to wait long enough before the next exciting pulse can be applied. Otherwise, the signal will be very low in the following acquisition. However, the most limiting characteristic is their high sensitivity to coherent macroscopic motions. As a consequence of the large duration and amplitude of the diffusion gradients all moving protons acquire an identical additional extra phase during each repetition time. Due to the properties of the Fourier transform, this extra phase does
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not lead to a higher attenuation of the signal, but to a spatial translation of the reconstructed image. During each repetition a different global phase is added and signal from the entire slice is received. Thus the image appears as a superposition of multiple similar images each one displaced parallel to the other by a different amount. The resulting ghosting artifacts which spread in phase-encoding direction make the image useless for diagnostic purposes (Fig. 2A). Different methods have been proposed to overcome the limitations which are imposed on the SE and STE sequences by motion artifacts. One of these is the navigator echo tech-
Fig. 2. An example of motion artifacts and the way of correction. (A) An image obtained from a volunteer with a diffusion-weighted SE sequence. Motion artifacts make this image non-diagnostic. (B) Motion artifacts are partially corrected using extra gating. Thus a reduced ghosting in the area of the cerebrospinal fluid can be observed. (C) The use of navigator echoes gives rise to a considerable reduction of artifacts in the upper part of the image. However, since this image is not triggered, artifacts at the height of the lateral ventricles persist. (D) A combination of both correction mechanisms result in an optimal image quality. Interestingly, a residual motion artifact caused by the eyes motions (see white arrow) can be also observed in image.
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nique, which was independently proposed by Ordidge et al. [11] and Anderson and Gore [12] in 1994 (see Fig. 1B). The basic idea is to measure and correct the motion-induced extra phases. With this aim an additional echo, known as navigator echo, is acquired without phase-encoding after each acquisition. In the absence of motion all the navigator echoes should be identical. Motion will affect in the same way both the phase-encoded echo and the navigator echo. Therefore, the differences between two successive navigator echoes allows one to calculate the relative extra phase added by the motion. By the subtraction of these phases most of the motion artifacts are suppressed (see Fig. 2C). Unfortunately, not all motion can be corrected to the same extent. For example, in DWI of the spine, cerebrospinal fluid pulsations still lead to artifacts after phase correction (see Fig. 2B). Therefore, an additional gating is usually needed in order to synchronize acquisitions with pulsations (see Fig. 2D). The combination of all these correction techniques allows to obtain almost artifact free images of the bone marrow [13]. However, such a procedure presents two important disadvantages. On the one hand, it is a technically complex method, and on the other hand, the acquisition time is considerably long. All these facts complicate the integration in the clinical routine. In order to increase the robustness of navigated diffusionweighted SE sequences, radial k-space SE sequences have been developed [14] (see Fig. 1C). This sequence type offers the advantage of self-navigation, since the center of the k-space is read in each line. Therefore, no additional echo for the phase correction is needed. Moreover, a selection algorithm, which discards the lines affected by complex motions such as regional and pulsating motions, reduces artifacts. The residual effects of cardiac motion and CSF pulsations also have to be suppressed by means of cardiac gating. This sequence has been demonstrated as a robust technique for high quality imaging of the spine [14]. Gudbjartsson et al. [15] have adopted a different approach in their line scan diffusion technique, which is a variant of the SE sequence (see Fig. 1D). They proposed to construct the image line by line from one-dimensionally acquired dataset, so that the relative phase between adjacent lines does not play any role. This sequence is inherently insensitive to motion and to susceptibility artifacts. In addition, due to a special selection scheme of the lines it is no longer necessary to wait for longitudinal relaxation before the next excitation, making this sequence considerably faster than the SE sequence. However, since the echo is produced only in a line and not in the whole slice, this sequence has a low SNR and therefore requires a careful optimization in order to keep an acceptable signal level. Another possibility of overcoming motion effects is to drastically decrease the acquisition time. MRI offers various fast acquisition techniques which can also be applied in DWI. Many of them are based on the successive refocusing of a diffusion-weighted echo, so that a complete image can be read after only one excitation. The intrinsic
transversal relaxation times, T2 or T2∗ , impose an unavoidable constraint to the time in which an echo can be refocused and, therefore, to the achievable spatial resolution of these sequences. 3.2. TSE-based diffusion weighted-sequences One way of refocusing the spins is by means of 180◦ pulses. The most relevant example of such a sequence is the turbo spin echo (TSE) sequence with the single-shot variants “rapid acquisition with relaxation enhancement” (RARE) [16] and “half-Fourier acquired single-shot turbo spin echo” (HASTE) [17]. These sequences have been adapted to DWI in spite of their sensitivity to the magnetization phase prior to the train of refocusing pulses, i.e. after its preparation. Small phase shifts of the magnetization with respect to the inversion axis of the 180◦ pulse lead to image artifacts. Therefore, a diffusion-weighted preparation of the magnetization, which inherently involves phase shift and spread, is expected to cause severe artifacts. However, appropriate modifications of these sequences allow to make them insensitive to the initial phase of the magnetization so that they become suitable for DWI. In order to modify the TSE-based sequences for DWI, a good understanding of the echo formation mechanism is needed. A central role plays the non-ideal character of the 180◦ refocusing pulse, which rotate the spins across the slice profile not exactly by 180◦ but by angles from 0 to 180◦ . Those spins which are not exactly inverted lead to stimulated and higher order echoes [18]. Therefore, the spins contributing to an echo, except from the first one, have been de-phased and re-phased in very different ways. Each one of these possible refocusing ways is known as a coherence pathway, and can be characterized according to the parity of the number of inversions that the spins have undergone. Spins following an even-parity pathway undergo an even number of inversions of the magnetization around the pulse direction, and thus the magnetization at the echo time keeps its initial orientation. However, spins in an odd-parity pathway suffer an odd number of inversions, thus leading to an inverted magnetization with respect to the inversion axis of the pulse. Since all spins contributing to an echo refocus at the same time, the component of the transversal magnetization which does not lie parallel to the axis of the refocusing pulse is progressively annihilated as the echo number increases and both types of coherence pathways turn to be equally populated. Two possibilities exist to avoid the destructive addition of coherent pathways. The first one is to split the even- and odd-parity echoes. A second solution is the elimination of the non-parallel component of the transversal magnetization prior to the first refocusing pulse. The “displaced ultra-fast low angle RARE sequence” (displaced U-FLARE) [19] belongs to the first group. In this sequence an extra de-phasing gradient is added before the readout gradient takes place (see Fig. 3A). Spins, which contribute to an echo following an even-parity pathway, add no net phase due to the new gradi-
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Fig. 3. Sequence schema of displaced U-FLARE, EPI and PSIF sequences. The diffusion gradients are always drawn with a diagonal hatch. (A) Displaced U-FLARE sequence. The part of the sequence outside the brackets prepares the transversal magnetization. The sequence part between brackets is responsible for the echo-refocusing, and is repeated until all the lines are acquired. The gradient in black is the de-phasing gradient introduced to split even- and odd-parity echoes. (B) Scheme of an echo-planar imaging sequence. As in A, the part of the sequence outside the brackets prepares the diffusion-weighting of the magnetization. Between the brackets is the part of the sequence that refocuses the echo. Each of the gradients in the readout direction allows the acquisition of an echo. After each echo a phase-encoding is performed by the small gradient in black (the so-called “blip”). (C) PSIF sequence. An excitation of a small angle, θ, is continuously repeated in order to achieve a steady-state of the magnetization. In each excitation a diffusion gradient is applied.
ent. On the contrary, all the spins in an odd-parity pathway acquire an extra phase. Thus echoes split in time according to their parity. Strong enough de-phasing gradients allow a shift of odd-echoes out of the acquisition window. A disadvantage of this procedure is presented by the signal intensity being reduced by a factor of two, since only the half of the echoes are acquired (see Fig. 4B for an example). Moreover, the maximum signal intensity is achieved after a few refocusing pulses, which implies that the first echoes must be discarded. Schick [18] proposed a sequence, “the split acquisition of fast spin echo signals for diffusion imaging” (SPLICE), which solve the SNR problem. In this sequence an unbalanced readout gradient is responsible for the splitting of the echoes, which are acquired separately. A final addition of the Fourier-transformed amplitude images of each echo results in an improvement of the SNR. A disadvantage of the SPLICE sequence is the increased echo spacing and thus the prolonged echo train duration.
Among the sequences in which the non-parallel component of the magnetization is eliminated, the “phaseinsensitive RARE pulse sequence” [20] introduced by Alsop stands out. Inserting a 90◦ pulse at one half echo time before the first refocusing pulse removes the out of phase component of the transversal magnetization which avoids the problem of destructive superposition of coherence pathways. Even more, the use of a tailored train of RF pulses produces a stable echo amplitude from the first echo and drastically diminishes the power deposition. The main disadvantage of this technique is that the signal is reduced by a factor of two due to the phase insensitive preparation of the magnetization. 3.3. Diffusion-weighted echo-planar imaging Another way of refocusing the spins is to use gradient echoes instead of spin echoes. The most relevant single-shot sequence by using a train of pure gradient echoes is the so-
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Fig. 4. Comparison of EPI, RARE and PSIF images obtained in the same patient. (A) 128 × 96 image acquired with an EPI sequence. As can be appreciated here, EPI images are normally degraded by artifacts, which seriously limit the clinical usefulness of the image. Some of the artifacts are caused by chemical shift (in this example, the structure indicated with arrow 1 may be a chemical shift artifact), and susceptibility inhomogeneities (see the distortion indicated with arrow 2). (B) 128 × 96 image acquired with a RARE sequence. Image quality is suitable for clinical purposes, though the image suffers from blurring. (C) 512 × 512 image interpolated from a 256 × 256 image acquired with a PSIF sequence. This high resolution image shows a sharp contrast and almost no artifacts. Thus PSIF images are very appropriate for clinical applications. The disadvantage of this sequence is that no quantification of the ADC can be performed.
called echo-planar imaging (EPI) sequence [21]. The adaptation of the spin echo EPI sequences to DWI requires a pair of diffusion gradients before the first acquisition (see Fig. 3B). Unlike the TSE sequence the extra addition of the diffusion gradients does not produce any undesired echo modulation and image artifacts. However, DW-EPI causes the same image artifacts as EPI [22], thus seriously restricting the applicability of DW-EPI for bone marrow imaging (see Fig. 4A). The most limiting artifacts are caused by off-resonance effects such as susceptibility artifacts. Common to all off-resonance effects is that not all the spins possess the same resonance (Larmor) frequency. To understand the implications of a resonance frequency spread for an EPI measurement let us consider an ideal experiment (see Fig. 5). Two spin species, A and B, with slightly different resonance frequencies are isotropically distributed in a cylindrical phantom, which lies parallel to a perfectly homogeneous external magnetic field. An ideal EPI experiment is performed on this phantom. First a 90◦ RF pulse turns over all the spins, which are included in a transversal slice located at one half of the phantom height. The flipped atoms of both species, initially in phase, start precessing. Since in the gradient echo technique the echo is produced without inverting the spins, differences in the resonance frequency produce a continuously increasing relative phase between both species, which otherwise re-phase at the same time (see in Fig. 5 the last two boxes). As it has been already seen, an additional phase will result in displacements in the reconstructed image. Therefore, the images produced from the species A and B will be displaced from each other (see the idealized imaging in Fig. 5). The actual way in which both images are displaced depends on
the accumulated phase out, i.e. on the way the k-space is sampled. In an EPI measurement the k-space is sampled in such a way that the points in readout direction are acquired very rapidly, whereas the sampling in the phase-encoding direction takes place more slowly, i.e. with a much lower effective receiver bandwidth (e.g. for a 128 × 128 acquisition matrix the sampling rate in phase-encoding direction is approximately by a factor of 128 greater than in the read out direction). Therefore, the accumulated phase difference between the two species is larger between two points in the phase-encoding direction than in the readout direction. Even more, since the phase difference grows at an uniform rate, the phase encoding is not altered and thus no smearing appears in the image. Obviously, the displacement caused by the accumulated phase difference must be compared to the pixel size in order to see whether visible artifacts occur. Normally, owing to the different acquisition rates in phase-encoding and readout directions, no artifacts can be seen in readout direction whereas, even by very small off-resonance effects (∼3 ppm), displacements in the phase-encoding direction are in the order of several pixels. Off-resonance effects are caused by distortions in the external magnetic field, chemical shift, or susceptibility inhomogeneities. With an adequate shimming it is possible to homogenize the external magnetic field in order to eliminate any artifact. Chemical shift can be reduced using a fat suppression or water-only excitation. This fact is of special importance in imaging of bone marrow due to the high content of fat within the yellow marrow. Imperfect fat suppression would lead to residual artifacts. The susceptibility in-
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homogeneities are the most problematic off-resonance effect when performing EPI in vivo, since a correction can hardly be performed. Susceptibility inhomogeneities mostly take place in the vicinity of interfaces between different media such as in the interfaces between bone, soft tissue, and air around the spine. Therefore, the applicability of DWEPI sequences to the bone marrow is seriously limited (see Fig. 4A). 3.4. Multi-shot echo-planar imaging A simple way to overcome the limitations that the T2∗ relaxation time imposes on the maximal echo train length and thereby, on the spatial resolution of EPI images is just to cover the complete k-space with different excitations and to combine all the data in a single raw data set [22]. Since each excitation requires a diffusion-weighting, patient motions can cause phase errors between the different data sets. Therefore, a navigated phase-correction is in the majority of cases necessary in order to prevent from motion artifacts. This technique has been proved to produce images of high diagnostic quality [22]. One of the major problems in breaking up the k-space is that phase and amplitude discontinuities between different data sets produce blurring and ghosting. Therefore, the most robust method to read out the k-space with a multi-shot EPI sequence is to acquire the k-space lines interleaved. With an adequate selection of the RF pulse flip angles a smooth variation of the signal intensities and phase throughout the k-space can be achieved avoiding artifacts [22]. Even more, this method is less sensitive to off-resonance effects. Since successively acquired lines are now separated by a number of lines equal to the number of excitations, Ne , the accumulated phase difference between spins in phase-encoding direction
Fig. 5. Ideal measurement with an EPI sequence in order to illustrate offresonance effects. A cylindrical phantom filled up with two spin species, A and B, with slightly different resonance frequencies lies parallel to a perfect homogeneous external magnetic field. An ideal EPI measurement, as depicted in the figure, is perform. In the boxes the behavior of the spins of type A and B in the rotating reference system (at the frequency of atoms of type A) is shown at different moments of the measurement. The spins are supposed to lie in the readout direction. Relaxation effects have not been considered in this idealized experiment. In the first box from the left the spins are represented just after the application of the 90◦ RF pulse. Both species are initially in phase. In the second box spins are represented while the de-phasing gradient is switched on. Because of the gradient in the readout direction, spins located at different positions on the readout direction accumulate a different phase. Due to their different resonance frequencies spins of type A and B have begun to de-phase (the white sector represents the phase difference between A and B, which is the same for all spins). Since the gradient echo does not turn over the spins, the phase difference continues growing through the measurement. Boxes three and four show how all the spins re-phase at the same time but with a growing phase difference (white sector). At the bottom the ideally resulting image is represented. The accumulated phase difference of B-spins causes a displacement of their image in the phase direction.
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is divided by Ne . Therefore, off-resonance effects in phase direction are reduced, thus making this sequence much more appropriate for DWI of the bone marrow than DW-EPI. 3.5. Steady-state free precession sequences A different type of fast acquisition sequences are those based on creating a steady-state of the longitudinal and transversal magnetization, the so-called steady-state free precession (SSFP) sequences. In these sequences the magnetization is repeatedly disturbed by a fast repetition of RF pulses. After some RF pulses the transversal magnetization
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achieves a steady-state, in which the amount of longitudinal and transversal magnetization after each pulse is the same. SSFP sequences can be adapted to DWI by introducing a diffusion gradient in each inter-pulse period (see Fig. 3C). The most useful of these sequences for the bone marrow is the diffusion-weighted PSIF sequence [23] (reversed fast imaging with steady precession (FISP [24]) sequence), which allows the acquisition of high resolution images in times of the order of few seconds because of the short repetition times of less than 50 ms. Since the RF pulse flip angle are always smaller than 180◦ , each pulse turns a part of the longitudinal magnetization into the transversal plane and partially brings the transversal magnetization back parallel to the magnetic field. Therefore, as in the TSE-based sequences, the echo is formed by the superposition of multiple coherence pathways. However, unless in the TSE-based sequences, every pathway has a different diffusion-weighting. The signal intensity depends on the sequence parameters and the diffusivity in such a very complicated way, that it is not suitable for quantitative determination of any physical constant [25]. Nevertheless, the good quality of the images showing almost no artifacts makes this sequence a very useful tool for diagnostics (see Fig. 4C).
4. Conclusions In this article we describe the different sequences that have been used for DWI of the bone marrow so far. The problems of each sequence type is analyzed carefully and their possible solutions are explained. Although the standard diffusionweighted SE and TSE pulse sequences are seriously affected by motion artifacts, images suitable for clinical use can be obtained either by adding a correction technique such as the navigator echo, or by using new developed SE-based sequence (such as the k-space spin echo or the line scan diffusion imaging sequences). Due to the bone marrow composition and location, single-shot sequences result in images which do not always offer the desired quality. For example, susceptibility artifacts of DW-EPI sequences seriously restrict the image quality. This problem is partially solved by the multi-shot DW-EPI sequence. SSFP sequences allow the acquisition of good quality images in short time, but with the disadvantage of not being appropriate for quantitative DWI. Unlike in the imaging of the brain, where DW-EPI has been established as the standard method for DWI, no generally accepted standard sequence for DWI of the bone marrow has been established.
Acknowledgment The authors would like to thank Dr. Sonja Buhmann for providing the MR images of Fig. 4 and for critical reading of the manuscript. This work was supported by
the Deutsche Forschungsgemeinschaft (DFG), Grant No. BA-2089(1-3).
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