M. Bock, PhD
Deutsches Krebsforschungszentrum (dkfz), Abt. Medizinische Physik in der Radiologie (E020), Im Neuenheimer Feld 280, 69120 Heidelberg, Germany
C O N T E N T S
11.1 Introduction 127
11.2 Gradient-Echo Techniques 128 11.2.1 Spoiled Gradient Echo 128 11.2.2 Fully Balanced Gradient Echo 131 11.3 Data Sampling 131
11.3.1 Partial-Fourier Techniques 131 11.3.2 View Sharing 132
11.3.3 Echo-Planar Imaging 133
11.3.4 Non-Cartesian k-Space Sampling 136 11.4 How to Combine
Fast Imaging Techniques 137
11.5 Real-Time Image Reconstruction 137 11.6 Conclusion 138
References 139
Fast Sequences for Dynamic and 11
Time-Resolved Imaging
Michael Bock
11.1
Introduction
Time-resolved MR imaging is increasingly used in the clinical routine as it allows for monitoring dynamic changes of the morphology and the organ function.
In general, the choice of the MR imaging technique is dictated by the time scale of the dynamic process to be imaged. If fl exing of the knee is to be visualised, conventional spin echo techniques can be used, since the motion of the knee can take place in between 2- min-long image acquisitions. The breathing motion,
if not forced, requires image acquisition times of the order of a second to capture the motility of the chest wall and the internal organs. The heart motion gives the most demanding time scale, where RR intervals of less than a second are found. In particular, the systolic motion of the heart and the blood in the adjacent ves- sels requires a temporal resolution of about 50 ms or less to suppress motion artefacts. All these given time scales are signifi cantly shorter when small children or infants are studied. In contrast agent studies, the time scale is furthermore infl uenced by the injection rate and volume as well as the injection site, so that a second-long acquisition time might be suffi cient to visualise contrast agent transit in the central arterial vasculature after injection through a peripheral vein, whereas sub-second time resolution is required for intra-arterial injection through a catheter.
A second, often confl icting requirement is given by the spatial dimensions of the target organ. To success- fully diagnose a disease, the spatial resolution needs to be adapted to the geometry of the target organ. In the example of knee imaging, this might not be problematic, since long acquisition times can be utilised and even small tendons can be visualised. During the breathing and the heart cycle, however, the temporal restriction of a second or a sub-second exists, and the delineation of a coronary arterial tree becomes diffi cult. In this case, use has been made of the fact that most organ move- ments are periodic. The MR image acquisition was segmented and synchronised with the organ motion using a physiologic monitor (e.g., ECG, breathing belt, and pulse oximeter). Later, the different segments were combined to form an artefact-free MR image series at different time points during the periodic organ motion.
Unfortunately, in contrast agent studies, signal changes are not periodic, and periodicity can also be destroyed by pathology (e.g., arrhythmia).
As has been explained in the previous sections, parallel imaging allows the signifi cant reduction of the image acquisition time over conventional MRI without any loss in spatial resolution. Thus, sub- second temporal resolution can be achieved even
with conventional fast gradient echo techniques. In this section, acquisition techniques will be presented that are typically used for dynamic MRI and that can be combined with parallel acquisition strategies. A more detailed description of fast imaging techniques and their implementation details can be found in textbooks (Vlaardingerbroek and den Boer 1996; Schmitt et al. 1998; Debatin and McKinnon 1998; Haacke et al. 1999; Bernstein et al. 2004).
11.2
Gradient-Echo Techniques
Today, fast MR imaging techniques predominantly use gradient echoes (GRE) for signal generation. In comparison to spin-echo (SE) techniques, GRE imag- ing offers several advantages for dynamic MRI with sub-second acquisition times:
In a SE pulse sequence, a 90°–180° RF pulse com- bination is employed to create an echo, whereas a GRE requires only a single RF pulse, so that the effective time for the acquisition of a line in k- space is signifi cantly shorter.
The lack of 180°-refocusing pulses in GRE MRI reduces the RF energy transmitted per unit time.
Thus, GRE techniques are less susceptible to limi- tations resulting from restrictions of the specifi c absorption rate (SAR).
The use of 90°-excitation pulses in SE MRI leads to a saturation of the longitudinal magnetisation.
After SE data acquisition, long signal recovery delays have to be introduced to establish suffi cient magnetisation, which unnecessarily prolongs the total scan time. In GRE imaging typically lower fl ip angles are used, and pulse sequences with very short repetition times still yield suffi cient signal.
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●
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In the following, two gradient echo pulse sequences are listed that are often used in dynamic studies.
Since naming conventions and acronyms differ from one vendor to the next, a list of the different pulse sequence names is given in Table 11.1.
11.2.1
Spoiled Gradient Echo
In the year 1985, Haase and co-workers proposed a pulse sequence that sampled a gradient echo using low fl ip angle RF excitation (1986). They called this pulse sequence Fast Low Angle SHot (FLASH) to distinguish their approach from the SE techniques existing at that time.
Following a slice-selective RF excitation with a fl ip angle D which is typically smaller than 90°, a fre- quency-encoded gradient echo is acquired (Fig. 11.1a).
After data acquisition, the remaining transverse mag- netisation is spoiled. Therefore, strong gradients are applied that create an intra-voxel dephasing before the next RF pulse. To increase the spoiling effect, the transmit phase of the RF pulses can also be varied in a pseudo-random fashion.
The repeated RF excitations at intervals TR lead to the establishment of a dynamic steady state of the longitudinal magnetisation. Assuming that the trans- verse magnetisation created by each RF excitation is no longer present at the end of each TR period (i.e., perfect spoiling), the signal equation for a FLASH pulse sequence can readily be computed:
S e
e e
( , ; , ) sin( )
cos( )
A R A
TR,TE T1 T2* A
TR/T1 TR/T1
TE/
1
1
T T2*
(11.1)
Here, U, T1 and T2* denote the object-specifi c spin density, longitudinal and apparent transverse relaxa- tion times, respectively, whereas D, TE and TR are the
Table 11.1
General Electric Philips Siemens Remarks
SPGR (spoiled GRASS) FFE-T1 FLASH
(fast low angle shot)
Spoiled gradient echo
GRASS (gradient-recalled acquisition in the steady state)
FFE
(fast fi eld echo)
FISP (fast imaging with steady precession)
Partially refocused gradient echo
FIESTA (fast imaging employ- ing steady-state acquisition)
Balanced FFE trueFISP Fully refocused gradient echo
SSFP
(steady-state free precession)
FFE-T2 PSIF
(time-reversed FISP)
Refocusing of spin echo coherence path in subsequent TR
Fig. 11.1a,b. Timing diagrams of a FLASH a and a trueFISP b pulse sequence. In both gradient echo pulse sequences, the slice-selective RF excitation is followed by a conventional phase and frequency encoding. In the spoiled gradient echo sequence, coherences of the transverse magnetisation are reduced by the application of strong spoiler gradients. In the trueFISP pulse sequence, all gradi- ent-induced phases are rewound (balanced) so that the transverse magnetisation fully contributes to the next RF excitation
b a
sequence-dependent parameters. In dynamic studies the parameters are often adjusted to create a specifi c contrast:
For very low fl ip angles of 10° and less, cos(D) is close to 1 and S is nearly independent of T1. If also TE is shorter than T2*, only spin density changes lead to a contrast in the FLASH image. Since spin density does not vary signifi cantly in many tissues (with the exception of lung tissue), images with low contrast are created (Fig. 11.2a).
●
If TE is of the order of the object’s T2*, a T2* con- trast is established in the images, if fl ip angles are kept low to avoid an additional interfering T1 con- trast. This contrast is advantageous in studies with a T2*-shortening contrast agent (e.g., ultra-small iron oxide particles, USPIO), where the arrival of the contrast in the target organ is seen as a signal reduction (Fig. 11.2b).
A T1-contrast is achieved by adjusting both TR and D simultaneously. For high fl ip angles D>60°
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●
Fig. 11.2a–d. Spoiled a-c and balanced d gradient-echo images with different parameter settings. a For very low fl ip angles (here, D=10°), a short TE of 3.6 ms and a moderate TR of 74 ms, a nearly constant signal is seen in the FLASH image (spin density weighting). b At a longer echo time (TE=12 ms), the differences in the T2* values lead to a T2* weighted image. c A short TE and a longer TR of 150 ms in combination with higher fl ip angles (D=70°) result in a T1-weighted FLASH image. d With a fully balanced gradient-echo technique (trueFISP) also the transverse signal components contribute to the steady-state signal and a mixed contrast is established, which pronounces tissues with a high T2/T1 ratio (e.g., fl uids). Here, the shortest possible TR and TE are chosen to avoid off-resonance artefacts
c a b
d
and a TR that is short compared to the object’s T1, a strong T1-contrast is established (Fig. 11.2c).
To avoid signal saturation at very short TRs of 10 ms and less, the fl ip angle is typically reduced to D=10°…30°, which still provides a moderate T1 dependence of the signal.
11.2.2
Fully Balanced Gradient Echo
In recent years, another GRE sequence has been found to have multiple uses in dynamic MRI: the GRE sequence with fully balanced gradient timing (Oppelt et al. 1986). In the spoiled GRE sequence, the transverse magnetisation was suppressed after data acquisition and the steady state was solely dependent on the longitudinal relaxation.
In the fully balanced GRE sequence, all gradients are rewound at the end of TR so that the transverse magnetisation also contributes to the steady-state signal. The fully balanced GRE sequence is often used in combination with an alternation of the transmit RF phase for every other RF pulse (Fig. 11.1b).
Compared to the spoiled GRE sequence, the fully balanced GRE sequence yields a signifi cantly higher signal since the transverse magnetisation is also uti- lised. Unfortunately, the image contrast is more dif- fi cult to compute and is dependent on the quotient of T2 and T1. Typically, this pronounces structures with a large T2/T1 ratio such as liquids (Fig. 11.2d). The fully balanced gradient timing refocuses the phase of the transverse magnetisation; however, so does every off-resonance. Therefore, these sequences are very sensitive to fi eld inhomogeneities that produce dark artefact bands in the image. Since the value of the off-resonance phase is proportional to the repeti- tion time, these artefacts are minimised by using the shortest possible TR.
11.3
Data Sampling
Gradient echo techniques provide a fast and effi cient way to acquire image data; however, at a repetition time of 5 ms and 256 phase encoding steps, the total acquisition time per image still amounts to 1.3 s, which is by far too long to resolve the dynamics of the beating heart. Several modifi cations of the image
acquisition have been proposed to accelerate the data acquisition process. Most of these techniques are independent of parallel imaging and can readily be combined with any parallel acquisition technique.
11.3.1
Partial-Fourier Techniques
In the ideal k-space representation of an MR image, a point symmetry relation (Hermitian symmetry) exists between data in the upper und the lower half of k-space (Fig. 11.3):
F
F; S S( ` a) ;( S`, Sa) * (11.2) Here, the asterisk denotes complex conjugation.
Essentially, this relation refl ects the fact that the phase of an ideal MR image should be constant everywhere in the image (i.e., without the loss of generality, it could be set to zero).
Fig. 11.3. In a partial Fourier k-space acquisition scheme, slightly more than half of k-space is acquired (grey areas), which results in an associated reduction in scan time. In the central region of k-space, a symmetric portion of the data is measured (shaded area) to estimate the low-order phase changes, which distort the Hermitian symmetry. When com- bined with parallel MRI, the original density of k-space lines is often maintained in this central section, because these data can also be used to calculate the weighting factors used for re- construction of the unmeasured lines in the k-space periphery (e.g., in the GRAPPA algorithm)
Unfortunately, in a realistic MR experiment, this relation only holds approximately, since there are local off-resonances that lead to a non-zero phase dis- tribution in the image. If one assumes that the phase varies only slowly over the image, a low-order phase estimate can be extracted from the central lines in k-space. This low-order phase estimate is then used to remove the phase variation and, subsequently, the lower half of k-space is reconstructed from data of the upper half. Techniques that sample only slightly more than one half of k-space are called partial-Fou- rier or half-Fourier techniques (Noll et al. 1991).
The assumption that the low-order phase estimate is suffi cient to correct for asymmetries in k-space is not valid at tissue boundaries in gradient echo images. Here, susceptibility changes introduce local static fi eld gradients, which do not infl uence spin echo images, but can lead to additional phase changes in gradient echo MRI. Since local phase changes are represented in the outer regions of k-space, partial- Fourier techniques often cannot recover the correct signal intensity at tissue boundaries.
Partial-Fourier imaging can favourably be com- bined with parallel imaging since both techniques require a central portion of k-space to be sampled at full k-space density (Bydder and Robson 2005). If, e.g., a 2/16th central section of k-space is acquired at full density and the remaining data in the upper half (7/16) at an acceleration factor of 2 (i.e., every other line is sampled), the total imaging time is reduced to (2/16)+(7/16)/2=11/32=34% of the original imag- ing time. During image reconstruction, the central k-space data are fi rst used to synthesise the missing lines in the upper half (e.g., using an autocalibrated algorithm such as GRAPPA), and then the partial Fourier reconstruction algorithm is applied.
Since both techniques – partial-Fourier and par- allel MRI – reduce the number of acquired k-space lines, an associated reduction in the SNR is observed.
With partial-Fourier techniques, however, the SNR reduction is constant over the image, since no local information is used during image reconstruction.
11.3.2 View Sharing
Another way to increase the temporal resolution at least nominally is called view sharing. In view shar- ing, portions of k-space are shared between images and, thus, need not be reacquired. Several concepts of view sharing have been proposed:
Keyhole MRI (van Vaals et al. 1993): In the most extreme version of view sharing, a full k-space data set is acquired only once (e.g., at the begin- ning of the dynamic series). Later, only the cen- tral portion of k-space is acquired and the missing outer k-space data are substituted using the lines from the full data set. Even though very high key- hole acceleration factors K (K=8 and more) can be achieved, the technique is not suitable in this form for clinical purposes, since the temporal variation in the high k-space regions (which is encoding the important edge features in the image) is not visualised at all (Fig. 11.4). In a variation of the keyhole technique, outer portions of k-space are sporadically re-acquired, which reduces the attain- able acceleration factor K. Keyhole techniques can readily be combined with parallel imaging using the full density reference data set(s) for coil cali- bration. For a parallel acceleration factor of R, only every R-th k-space line is acquired in the dynamic phase. Depending on the frequency of repetition of the reference data set, an accelera- tion factor of nearly K×R can thus be achieved during the dynamic phase using parallel acquisi- tion techniques. Sampling of the coil calibration data is preferably integrated in the acquisition of the reference data set(s) to achieve a maximum acceleration during the dynamic phase.
TRICKS (Korosec et al. 1996): The time-resolved interpolated contrast kinetics or TRICKS technique is similar to keyhole imaging in design, though here k-space is not only divided into a central and an outer region, but several sections, A, B, C, etc., where the letters denote an increasing distance to the k-space centre. In TRICKS notation, the key- hole technique as described above acquires a full data set (A-B) followed by the repeated acquisi- tion of the central section A: (A-B)-A-A-A-A-…
With TRICKS, the central regions are acquired more often than the outer k-space regions; how- ever, all regions of k-space are updated continu- ously (though not with the same frequency). For a segmentation of k-space into four regions A, B, C, and D, a TRICKS data acquisition could look like this: (A-B-C-D)-(A-B)-(A-C)-(A-B)-(A-D)-(A-B)- (A-C)-(A-B)-(A-D)-… A k-space data set would be synthesized from these data for each central section A using the nearest peripheral sections B, C, and D in time. All algorithms for parallel MRI can be used in combination with TRICKS; how- ever, the highest acceleration is achieved when the coil sensitivity information is sampled fi rst (e.g.,
●
●
during the initial A-B-C-D part). In the dynamic phase, all data are acquired with full parallel acqui- sition acceleration and without any autocalibra- tion data, thus increasing the temporal resolution.
The combination of parallel imaging and TRICKS for fast angiographic MRI has also been referred to as the time-resolved echo-shared angiography technique, or TREAT.
View sharing during ECG-gated scans (Foo et al. 1995): A special form of view sharing can be found in gated acquisition techniques (e.g., during ECG gating), where the total acquisition time is reduced by acquiring not only one but several k- space lines per synchronisation cycle and by shar- ing some of the data between successive (cardiac) phases. Compared to the previous techniques, ECG-synchronised view sharing is slightly differ- ent, since not all of k-space is acquired in one car- diac cycle and the acquisition is repeated several times. In the TRICKS notation ECG-gated view sharing with three k-space regions could, e.g., be realised in the following way: (ECG trigger)-Ai- Bi-Ai-Ci-Ai-Bi-Ai-Ci-…(ECG trigger) -Ai+1-B i+1-A
i+1-C i+1-A i+1-B i+1-A i+1-C i+1-… Here, the index i
●
denotes the ith k-space line. View sharing in the cardiac cycle leads to temporal blurring, since the reconstructed temporal resolution is higher than the measured one: in the above example, in every second TR, an A-segment is sampled and the reconstructed time difference is 2·TR; however, 3·TR are required to fi ll k-space.
11.3.3
Echo-Planar Imaging
In conventional MRI pulse sequences, one line of k- space is acquired per RF excitation. This image acqui- sition technique is time-ineffi cient, since a constant fraction of the total scan time (50% and more) is spent on magnetisation preparation and is not used for data acquisition. In the year 1977, Sir Peter Mansfi eld proposed a method to acquire all k-space lines follow- ing a single RF excitation, which he called echo planar imaging Mansfi eld (1977). Essentially, the gradient echo in readout direction is refocused using gradients of alternating polarity, while a constant weak gradient or small blipped gradients are present in phase-encod-
Fig. 11.4a–b. Initial full-resolution reference image a and subsequent keyhole image b, where only the central 1/8th of k-space was re-acquired and the k-space periphery was taken from a. In areas where motion has occurred such as the tongue and the larynx artefacts can be seen (arrows) and the precise location of the anatomical structures cannot be recovered
b a
ing direction (Fig. 11.5). The combination of these gra- dients leads to meandering k-space trajectory, and, if combined with an appropriate dephasing, allows to sample k-space in less than 100 ms.
Echo planar imaging (EPI) became clinically available in the early 1990s when powerful gradi- ent systems (Gmax>20 mT/m) with high slew rates [smax>40 T/(m·s)] were introduced. The time-effi - ciency of an EPI pulse sequence can easily exceed 80%. Unfortunately, single-shot EPI pulse sequences suffer from several artefacts that limit their applica- bility in clinical MRI. During the long readout gra- dient echo trains, the signal decays with the appar- ent transverse relaxation time T2*, which creates a fi lter effect on the k-space data and limits the spatial resolution. Image distortions are often seen at tissue boundaries where the susceptibility changes lead to static fi eld gradients that are of the same order of magnitude as the weak phase-encoding gradients.
The long echo trains also make the data acquisition prone to off-resonance artefacts, which are often suppressed using water-selective excitation pulses.
Single-shot echo planar images still suffer from arte- facts in anatomical regions with heterogeneous tis- sues such as the abdomen.
A compromise between sampling effi ciency and image quality is given by the segmented EPI tech- nique, where only a certain number of k-space lines
are acquired per RF excitation (McKinnon 1993).
This k-space segmentation requires repeated RF exci- tations and a combination of k-space lines from dif- ferent excitations, which makes segmented EPI more susceptible to motion artefacts. With segmented EPI the echo train length can be reduced to a level where distortion artefacts are tolerable and image acquisition times are still considerably shorter than conventional gradient warp techniques (Fig. 11.6).
In general, the acquisition time per image TA can be written as:
TA lines RF
echoes
¥ ADC
§¦ ´
¶µ
N t
n t (11.3)
where tRF denotes the time needed for RF excitation and initial phase encoding, tADC is the time required to sample a single line in k-space, and nechoes is the number of gradient echoes per RF excitation. In Fig. 11.7, the acquisition time for Nlines=77 is shown as a function of nechoes.
In echo planar imaging, additional reference scans are often performed to compensate for hardware imperfections and off-resonance effects. The com- bination of data acquired with positive and negative readout gradients often introduces small systematic differences between odd and even k-space lines, which manifest in ghost images shifted by half the fi eld of view (N/2 ghosts or Nyquist ghosts). In the
Fig. 11.5. Gradient tim- ing of a segmented echo planar imaging pulse sequence. Following a conventional slice-selec- tive RF excitation, a series of gradient echoes is acquired (here: four). In phase-encoding direc- tion, a variable dephaser gradient table is used to shift the start of the k-space trajectory from the k-space centre to the periphery, and small trapezoidal gradients (so- called blips) are utilised for advancing to encode the different k-space segments. In the most extreme case, the number of echoes is equal to the matrix size, which results in the standard single- shot EPI pulse sequence
Fig. 11.6. Segmented EPI images of the heart with necho=3, 5, 7 (top row) and 15, 29, 45 (bottom row) lines per TR. Due to the linear k-space encoding, TE increases with the number of lines from 2.9 ms to 16 ms, which leads to an increased sensitivity to motion artefacts, e.g., in the heart chambers. With increasing necho also the fat-water shift increases (arrows) so that fat satura- tion techniques become necessary
Fig. 11.7. Image acquisition time for a seg- mented EPI pulse sequence as a function of the number of echoes acquired per TR.
The repetition time TR is minimized for each sequence and therefore increases with the number of echoes per TR. The points are values taken from a measure- ment, the solid line is the theoretical prediction and the dotted line takes an additional repetition for the measure- ment of reference lines into account. In a conventional data acquisition, only a single line is acquired per TR, whereas all lines (here: 77) are sampled in one TR a single-shot EPI. Segmented acquisition schemes allow trading acquisition time for echo train length so that both higher matrix sizes and lower artefact sizes are possible
reference scans, which are acquired without phase encoding, these shifts can be measured and used to correct the k-space lines before reconstruction.
In segmented EPI acquisitions, larger portions of k-space are acquired with the same echo time, and phase discontinuities can occur between subsequent k-space segments. Here, the technique of echo time shifting (McKinnon 1993) can be employed to mini-
mize phase discontinuities, however, at the expense of slightly prolonged repetition times.
With parallel imaging, fewer k-space lines need to be acquired, which allows reducing the echo train length and, accordingly, the blurring in phase-encod- ing direction. Additionally, the step width in k-space is increased by the use of stronger phase-encoding gradients, which helps to reduce the distortion arte-
facts. Even though the acquisition of coil sensitivity information can in principle be integrated into every EPI echo train, again the highest acceleration factors are achieved when extra reference scans with sepa- rate RF excitations are acquired prior to the dynamic phase of the acquisition. For acceleration factors of 3 and higher, these scans are typically segmented to achieve a similar timing as during dynamic data acquisition.
11.3.4
Non-Cartesian k-Space Sampling
As has been outlined in Chapter 6, MR raw data can be acquired on a non-Cartesian grid in k-space. In particular, radial (Lauterbur 1973) and spiral (Ahn et al. 1986) acquisition techniques have been used
in the past for both 2D and 3D acquisitions. These special variants of non-Cartesian data sampling have several unique properties that render them well suited for fast MRI applications:
Both imaging techniques start acquiring data at the centre of k-space. Since no additional spatial encoding is required between RF excitation and the beginning of the data acquisition, both meth- ods can achieve very short echo times (TE<1 ms).
At short echo times, motion artefacts are mini- mised, and tissues with very short T2* relaxation times (e.g., lung tissue with T2*<2 ms at 1.5 T) can be visualised.
In particular during radial scanning, the centre of k-space is sampled more often than the periph- ery. During data reconstruction, this leads to an averaging of the low frequency components and additionally reduces motion artefacts (Fig. 11.8).
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Fig. 11.8. Radial MR image time series showing the uncooperative author who is speaking and violently shaking his head. Due to the oversampling of the k-space centre motion artefacts are averaged, and no ghost images of the moving structures are seen. Also in this example view shar- ing was used to increase the temporal resolution
To fulfi l the Nyquist criterion, S-times more radial lines need to be sampled than rectilinear k-space lines. This number can be reduced to S/2, if the radial acquisition is not started at the k-space origin, but on one side of k-space, and the radial line is traversing the k-space centre. For fast MR applications, angular undersampling is employed and the number of radial lines is reduced sig- nifi cantly. Even though streak artefacts are thus induced in the MR images, a high spatial resolu- tion (i.e., the resolution along the radial line) can be maintained in the reconstructed images.
Both radial and spiral MRIs are often used in com- bination with view sharing. Therefore, a number n of the total number of lines (or spirals) N is re- acquired after an initial full k-space data set, and the data are dynamically exchanged in k-space before image reconstruction.
Even though combining parallel imaging with radial or spiral MR techniques is not as straightfor- ward as with conventional Cartesian k-space sam- pling, the concepts are essentially similar (Pruess- mann et al. 2001). Using low-resolution coil sensitivity information, missing radial spokes or spiral arms are reconstructed as a weighted linear combination of the measured k-space data, where the weights are determined from the coil sensitivities.
Radial data sampling covers k-space more densely in the centre than in the periphery, which is not neces- sarily true for spiral MRI, where the distance between individual spiral arms can be much larger than the distance between data points along the spiral. The radial data oversampling near the k-space centre can be exploited during parallel imaging, because thus radial MRI automatically provides a low-resolution image of the coil sensitivities during image acquisi- tion (Yeh et al. 2005).
11.4
How to Combine Fast Imaging Techniques In the previous sections, several techniques to accel- erate the MR image acquisition have been described.
All of these techniques can be combined with parallel imaging, which results in an acceleration of the image acquisition. Furthermore, a mixture of techniques is possible, resulting in an even faster image acquisi- tion scheme.
●
●
For a conventional trueFISP pulse sequence with TR=4 ms and 256 phase encoding steps, the acquisi- tion time per image is about 1 s. Using a 6/8th par- tial Fourier technique, this acquisition time can be reduced to 768 ms, and the use of an acceleration factor of 3 in parallel imaging leads to an acquisition time of about 260 ms. A further reduction could be achieved with a segmented EPI acquisition, where, e.g., three lines are acquired per TR. Under typical conditions this could reduce the acquisition time by 40%, so that the fi nal image acquisition time per image is 160 ms. Unfortunately, the prolonged TR would also lead to an increased susceptibility to off- resonance artefacts so that this technique might only be applicable in anatomical regions with a high B0 fi eld homogeneity such as the brain.
Another possible combination is the use of a spoiled gradient echo technique with a radial data acquisition to visualise the blood vessels after infu- sion of a T1-shortening contrast agent. Using angu- lar undersampling, the number of radial spokes can be kept signifi cantly shorter than the number of k-space lines required for a Cartesian data acquisi- tion. At a TR of 4 ms, Nradial=100 radial spokes, the image data can be acquired in only 400 ms; however, streak artefacts from angular undersampling are to be expected in the fi nal MR images. With a parallel imaging technique the acquisition time can easily be halved (TA=200 ms), and view sharing between sub- sequent radial data acquisitions could be used to fur- ther increase the (nominal) temporal resolution to 100 ms and less. Again, the high temporal resolution is traded against artefacts from temporal averaging and spatial undersampling.
11.5
Real-Time Image Reconstruction
In Cartesian MRI, image data are acquired on a rectilinear k-space grid and a simple Fourier trans- form (implemented with the fast Fourier transform algorithm, FFT) is suffi cient to reconstruct the MR image. Even though the FFT needs to be performed separately for the data from each coil element, images can be reconstructed in (nearly) real-time using con- ventional computer hardware.
Already the introduction of a partial Fourier acquisition scheme increases the load on the compu- ter system, because more images need to be recon-
structed in the same time. Furthermore, the recon- struction is more complicated and time-consuming, because the low-order phase estimate needs to be reconstructed and subtracted. View sharing addi- tionally increases the computational burden, since even more images are reconstructed per unit time.
Non-Cartesian image acquisition techniques are even more time-consuming during image reconstruction, because raw data are typically fi rst interpolated to a Cartesian grid, and k-space density compensation needs to be performed.
The increased temporal resolution in the recon- structed images is thus often achieved through sig- nifi cantly prolonged image reconstruction times of several minutes. Even though ever-faster computer hard- and software are available, often image recon- struction cannot keep up with the high data rates during parallel imaging. Nevertheless, parallel acqui- sition techniques were successfully combined with real-time MRI to monitor interventional procedures.
In Fig. 11.9, an active catheter tracking experiment is shown, where the slice position was automatically adjusted to the position of a small tracking coil at the tip of a catheter. With a trueFISP pulse sequence, a
Fig. 11.9. Three images out of a real-time trueFISP time series acquired with active catheter tracking. In between the acquisi- tion of two subsequent time frames, the position of a small tracking coil (arrows) was determined, and the imaging slice was automatically shifted to the current coil position. To accelerate the data acquisition, parallel imaging was employed (parallel imaging factor 2) so that the acquisition time per image could be reduced to 210 ms. At this high temporal resolution, mo- tion artefacts of the beating heart of the animal were hardly visible. Since the image position was continuously changing, an algorithm with auto-calibration (GRAPPA) was chosen. With fi ve imaging coils, a real-time image reconstruction was possible using an optimised reconstruction algorithm
temporal resolution of 5 images/s is achieved using a GRAPPA factor of 2. Note that during interventional procedures the reconstruction factors for parallel imaging need to be re-acquired for each slice posi- tion. This limits the choice of parallel acquisition techniques to auto-correlation methods, where refer- ence lines are continuously re-acquired for each new image position. To accelerate the reconstruction, all those imaging coils that were too far away from the imaging slice were dynamically excluded from the reconstruction (Müller et al. 2005).
11.6 Conclusion
Parallel-imaging techniques can readily be com- bined with nearly every existing fast MRI method.
Coil sensitivity information is preferably acquired before the dynamic scan starts to increase the tem- poral resolution. When combined with partial-Fou- rier acquisition strategies, coil sensitivities can also
be extracted from the same data as are required for low-order phase correction. Real-time parallel MRI is feasible, though challenging, because fast parallel- image reconstruction requires powerful computers and optimised reconstruction algorithms.
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