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Keywords:

  • microscopy;
  • steady-state;
  • banding;
  • artifact

Abstract

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES

High-resolution imaging of trabecular bone aimed at analyzing the bone's microarchitecture is preferably performed with spin-echo-type pulse sequences. Unlike gradient echoes, spin-echoes are immune to artifactual broadening of trabeculae caused by local static field gradients near the bone–bone marrow interface and signal loss from chemical shift dephasing at k-space center. However, the previously practiced 3D fast large-angle spin-echo (FLASE) pulse sequence was found to be prone to a low-frequency modulation artifact in both the readout and slice direction. The artifact is caused by deviations in the effective flip angle of the nonselective 180° pulse, which converts a fraction of the phase-encoded transverse magnetization to longitudinal magnetization. The latter recurs as transverse magnetization in the subsequent pulse sequence cycle forming a spurious stimulated echo. The objective of this work was to perform a k-space analysis of this steady-state artifact and propose two modifications of the original 3D FLASE that effectively remove it. The results of the simulations were in exact agreement with the experiments and the proposed remedy was found to eliminate the artifact. Magn Reson Med 52:346–353, 2004. © 2004 Wiley-Liss, Inc.

There is growing evidence that the mechanical competence of trabecular bone (the type of bone in which the majority of osteoporotic fractures occur), and thus its resistance to fracture, is significantly determined by the bone's architectural make-up (1–5). In vivo MR microimaging has proven its potential for assessing the bone's 3D structure. Nevertheless, obtaining reproducible structural parameters strongly depends on achieving artifact-free images in a resolution regime on the order of 100–150 μm (6).

The magnetic properties of bone and the chemical composition of bone marrow pose stringent requirements on pulse sequences appropriate for MR microimaging of trabecular bone. Strong susceptibility-induced gradients as well as the multiple spectral components of triacyl glycerides in fatty bone marrow cause dephasing that diminishes the marrow signal. At the interface with the trabeculae the susceptibility-induced local gradients lead to signal loss that causes the trabeculae to appear artifactually thickened in the reconstructed image (7). In a spin-echo sequence, the signal loss caused by local susceptibility-induced gradients is fully recovered and the isochromats from the various spectral components of fat are rephased at k-space center. These two reasons make spin-echo type sequences particularly well suited for imaging trabecular bone. However, patient comfort and the large number of pulse sequence cycles inherent to 3D imaging put limits on the allowable repetition time, TR.

Two previous publications from this laboratory introduced 3D FLASE (fast large-angle spin echo), a pulse sequence optimized for the purpose of imaging trabecular bone (8, 9). The pulse sequence has been used in prior trabecular bone imaging studies (5, 10) and is an integral element of the virtual bone biopsy (6). 3D FLASE is a large-flip-angle, 3D spin-echo sequence, shown in Fig. 1a). The fundamental principle of FLASE is the dual function of the 180° pulse that serves both for phase-reversal and for restoration of the inverted longitudinal magnetization created by the initial excitation pulse. A slice-selective excitation with a flip angle of 140°, used in 3D FLASE gives maximum signal for a repetition time of 80 ms based on estimated bone marrow T1 of 300 ms and T2 of 60 ms (8). The excitation pulse is a minimum-phase Shinnar-Le Roux (SLR) pulse of 3.2 ms length, while the refocusing pulse is a hard 180° pulse of 0.5 ms. Crusher gradients in the slice-encoding direction are applied before and after the refocusing pulse to dephase the transverse magnetization produced by the imperfection of the refocusing pulse. The final pulse played out is a killer gradient in the slice-encoding direction. This gradient has the effect of averaging the signal within one voxel over different steady-state magnetizations. Its magnitude causes the phase within one voxel to be spread over several cycles (about 10) along the slice direction. In this manner any sensitivity of the steady-state magnetization to local field inhomogeneities is removed as the various magnetization components are averaged out in each voxel. The readout bandwidth is 16 kHz and a fractional echo sampling 312 data points is typically acquired. Between the rewinder pulses and the killer pulse (shown in Fig. 1a), a navigator echo (9, 11) is acquired for retrospective motion correction (not shown on Fig. 1a). Here, we shall ignore this part of the sequence since it does not have RF pulses associated with it and its gradients are fully balanced and of no consequence to the steady-state evolution of the magnetization, which is the subject of the current analysis.

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Figure 1. a: 3D FLASE pulse sequence. b: Axial slice from a 32-slice 3D FLASE scan of the wrist acquired at 1.5 T (General Electric Signa, Milwaukee, WI) at a voxel size of 137 × 137 × 410 μm3 using an FOV of 7 × 4 cm and 512 × 288 × 32 matrix. Note banding artifact in readout direction (left to right); banding modulation is also present in the slice-encoding direction as evidenced in the coronal (bottom) and sagittal reformations (left).

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A significant number of scans obtained using 3D FLASE display a low-frequency modulation in the readout and slice directions, an extreme case of which is shown in Fig. 1b. The spatial frequency of the banding is found to be resolution-dependent, the period of the banding being larger at lower resolution. The artifact is also insensitive to the strength of the killer gradients. Further, the intensity of the artifact varies from scan to scan and on some occasions degrades the image quality to the extent that repeat scans are needed. In the following we describe the origin of this artifact and discuss different approaches toward its removal.

THEORY AND METHODS

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES

Even though spin-echo sequences are well suited for imaging of trabecular bone, they are sensitive to the quality of the inversion pulse. The flip angle of the refocusing hard pulse may deviate from 180°, depending on the actual amplitude of the applied RF field.

In the case of a dedicated wrist transmit/receive coil, the problem of flip angle calibration is specific to the coil geometry rather than the particular scanner implementation. The custom-built, elliptical cross-section quadrature birdcage coil (12) used in our laboratory has been designed for optimum filling factor. It has high RF field homogeneity in most of its central region. However, near the coil conductors the field is greater and there is a transition region where it is relatively inhomogeneous. Depending on wrist size, the subcutaneous fat is located in this transition region. Since the flip angle in the transition region differs from that near the coil axis, automatic setting of the transmit gain on the basis of the total detected spectral signal (which is dominated by the signal from fat) may therefore fail to produce the correct flip angle in the volume of interest. Thus, although automatic optimization often gives good estimates of the transmit gain, the operator manually checks each estimate. Manual transmit gain adjustment has the advantage that it is based on a projection image, therefore allowing flip angle optimization in the volume of interest. Even though the projection contains signal from the proximity of the coil edges the central part of the projection, used for adjustment, is dominated by signal from the central region of the coil. Nonetheless, even with these precautions, slight misadjustments are usually unavoidable.

While these imperfections are generally small, in combination with the particularities of our sequence they can produce two noticeable artifacts. The first is a consequence of longitudinal magnetization being rotated into the transverse plane after the imperfect refocusing pulse. This signal is dephased by the trailing crusher gradient applied immediately after the refocusing pulse. The remaining unwanted signal is then “chopped” by alternating the phase of the refocusing pulse between 0° and 180°. In this manner, the unwanted signal is displaced to the end slices of the reconstructed image, which are subsequently discarded (8).

Another consequence of the imperfection of the phase-reversal pulse is that a portion of the transverse magnetization, which is phase-encoded before the refocusing pulse, is converted into longitudinal magnetization (coherence pathways II and III in Fig. 2). Because of the short repetition time (TR = 80 ms) relative to the longitudinal relaxation time (T1 = 300 ms in fatty bone marrow), the longitudinal magnetization created in this manner is not significantly attenuated by the end of the repetition period and is rotated back into the transverse plane by the excitation pulse of the subsequent pulse sequence cycle. It is then again phase-encoded and refocused by the 180° pulse at time TE/2 after the main echo, as shown in Fig. 2. This unwanted echo, labeled II in Fig. 2, is the source of the described banding artifact.

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Figure 2. Coherence pathways for the 3D FLASE pulse sequence for two consecutive excitations in the presence of imperfections in the 180° pulse. In the subsequent pulse sequence cycle, at time t4, two echoes (generated via pathways labeled I and III) contribute to the signal at the center of k-space and a single stimulated echo at time t5 (generated via pathway II) produces the artifactual signal TE/2 after the main echo. Note that the artifactual echo is not affected by the crusher gradient following the refocusing pulse, since at that time the magnetization for that pathway is longitudinal.

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Coherence Pathway of Artifactual Signal

Inspection of the coherence pathways of the signals generating the artifact (pathway II in Fig. 2) and the main echo (pathway I in Fig. 2) clarifies the origin of the artifactual signal and allows characterization of its relationship to the main echo. We will only consider coherence pathways I, II, and III, since only effects stemming from coherence pathways connecting two adjacent repetition periods are relevant. There are three more coherence pathways that can produce echoes in the following TR period (pathways IV, V, and VI in Fig. 2). In all of these coherence pathways the magnetization is transverse, up to the point it forms its strongest echo in the next repetition period. Therefore, the magnetization amplitude is significantly attenuated by T2 decay (T2 = 60 ms, TR = 80 ms) such that its amplitude relative to that of the magnetization in coherence pathway II is negligible.

For simplicity, we assume that all phase encodings and readout dephasing occur before the refocusing pulse, as shown in Fig. 1a. It is noted that in the implementation used for imaging at very high resolution, the y-phase encoding is split such that some of the encoding is performed after the refocusing pulse (8) in order to shorten the echo time. However, the essence of our arguments does not change for this particular variant of FLASE.

Even though all gradients varying between excitations are balanced within one repetition period, the magnetization is never fully in the steady state. The gradients are only balanced if one considers all zeroth moments within a TR period. It is evident, however, that the moments are not fully balanced in the periods between the excitation and the refocusing pulses. As already indicated, an imperfect refocusing pulse converts the phase-encoded transverse magnetization into longitudinal magnetization with phase memory that causes leakage (13, 14) into the following repetition period. However, since this effect is small we shall assume that the difference in the initial magnetization for two adjacent repetitions is negligible, which will simplify the analysis.

To propagate the amplitude of the magnetization of interest we use a standard form for the change in magnetization after a rotation by angle an θ about the x axis at time t (15):

  • equation image(1)
  • equation image(2)

where, M+(t+) (M+(t−)) is the complex transverse magnetization just after (before) time t and, analogously, Mz(t+) (Mz(t−)) is the real amplitude of the longitudinal magnetization immediately after (before) time t, while * denotes complex conjugation. We first follow coherence pathway II marked by the bold line in Fig. 2. The angle of the SLR pulse will be denoted by θ and the deviation of the refocusing pulse flip angle from π will be denoted by Δϕ. We set the amplitude of the transverse magnetization just after the SLR pulse to 1, since we are interested only in relative amplitudes. Under this assumption the transverse magnetization has the following amplitude just before the refocusing pulse

  • equation image(3)

where ϕ is the phase imparted by phase encoding and readout prephasing, and ϕc is the phase due to the crusher prior to the phase-reversal pulse. All times referred to further in the text are labeled in Fig. 2. The part of the longitudinal magnetization that is produced by the refocusing pulse in coherence pathway II has the following magnitude and phase at the end of the first repetition period

  • equation image(4)

The phase imparted by the crusher gradient is now stored in the longitudinal magnetization and is not affected by the trailing crusher gradient. After the excitation pulse in the next repetition this magnetization is rotated back into the transverse plane and phase encoded again

  • equation image(5)

The artifactual echo is formed after the refocusing pulse with amplitude

  • equation image(6)

at a point in k-space determined by the phase factor in Eq. [5] and the phase rewound by the trailing crusher. After this refocusing pulse the artifactual magnetization is in the transverse plane and the phase accumulated from the application of the left crusher gradient in this repetition is rephased by the trailing crusher; however, the phase encoding imparted by the leading crusher in the previous repetition persists.

The amplitude of the main echo originating from coherence pathway I is

  • equation image(7)

Note that the T2 decay is less for the main echo since the magnetization generating it spends half as much time in the transverse plane as the magnetization forming the artifactual echo. The magnitude, ϵ, of the latter relative to the main echo is the absolute value of the ratio of expression [6] over [7]:

  • equation image(8)

Using values for the parameters given earlier and TE = 9 ms, we infer that the relative magnitude of the artifact is given as:

  • equation image(9)

Using Eq. [9] and measuring the height of the main and artifact echoes in the raw k-space data, we estimate an average imperfection angle of 25° for the extreme case shown in Fig. 1b. In clinical scans banding stemming from imperfection angles of about 4° is noticeable, while pronounced banding occurs with imperfection angles of about 10°. These angles correspond to relative amplitude values of ε = 1.5% and ε = 3.8%.

We also note that coherence pathway III shown in Fig. 2 produces an echo simultaneous with the primary echo, thus increasing the amplitude of the main signal.

k-Space Analysis of Image Artifact

From Eq. [5] we see that the artifactual signal was phase encoded twice in y and z and also prephased twice in x. Thus, the nonbalanced phase imparted by the leading crusher before the first refocusing pulse moves the center of the echo along the kz axis by Δkz. In the readout direction the artifactual spin echo will appear TE/2 later than the main echo (t5 in Fig. 2) since this is the time needed for the stimulated echo to occur (see Fig. 2). The artifactual gradient echo in the readout direction occurs slightly earlier than t5 because the readout does not start immediately after the refocusing pulse due to the trailing crusher. In this manner, the center of the artifactual echo is translated by kpre in the kx direction. The magnitude of kpre is determined by the fractional echo dephasing and sets the spatial period of the banding in the readout direction, as will be shown further in the text. The artifactual signal sa can be written as:

  • equation image(10)

where s is the pristine signal of the main echo and ε is the amplitude of the artifactual echo relative to the amplitude of the main echo. The total signal stotal is then given by:

  • equation image(11)

To analyze the effect of the spurious echo on the reconstructed magnitude image, we calculate the magnitude of the corrupted image using Eq. [10]. We first calculate the Fourier transform of the total signal ρtotal(x,y,z) = F[stotal(kx,ky,kz)] that gives the following complex image:

  • equation image(12)

where ρ = F[s] is the image that would be obtained from the pristine data s. The magnitude image is:

  • equation image(13)

where we have assumed that ρ(x,y,z) is real. The first term under the square root of Eq. [13] is the pristine image. The next two terms describe the image degradation due to the artifact. It would be desirable to expand Eq. [13] into a Taylor series since this would produce additive corruption terms. Even though ε is small, we cannot do this for all points. The ratio of the “stretched out” image ρ(x,y/2,z/2) and the pristine image can become much larger than ε if ρ(x,y,z) is very small in some region of the image. Because of this we shall consider only the full expression and resort to numerical calculations to compare with the experimental results.

The term of Eq. [13] that is linear in ε produces the main banding artifact observed in clinical scans. This term becomes noticeable only where the pristine image has sufficiently high intensity since it scales as ρ(x,y,z). It is also modulated by the stretched-out image ρ(x,y/2,z/2) and the cosine banding factor. In clinical scans the modulation from the two image terms appears as inhomogeneities in the cosine banding. Due to the small magnitude of the artifact signal and complicated shape of the image modulation, other details are not discernable. The banding modulation occurs in both readout and slice directions, as shown in Fig. 1b. This was observed in clinical scans as a translation of the banding along the readout direction as one would cycle through slices. Also, for a fixed field of view kpre is smaller for lower readout resolutions, leading to a longer period of the banding in the readout direction, as observed. Finally, the third term of Eq. [13] is quadratic in ε and appears as a faint ghost of the pristine image, stretched out by a factor of two in the y and z phase-encoding directions. One should also note that, since, as mentioned before, the stimulated secondary echo at t5 is not exactly coincident with the gradient echo, the artifactual signal may carry some extraneous phase (from background gradients) relative to the main signal. This phase can lead to curvature of the banding, as seen in the upper portion of Fig. 1b.

RESULTS AND DISCUSSION

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES

Figure 3a displays a 3D FLASE scan of a trabecular bone specimen acquired with suboptimal transmitter gain setting to enhance the artifact. The phantom was a 1-cm section of demarrowed human bone placed in a cylindrical container of 5 cm diameter filled with formalin doped with Gd-DTPA. Figure 3b shows an image simulated using Eq. [13] assuming that the pristine image was of a homogenous cylinder parallel to the z-axis and spanning from one to the other end of the image volume. Since the pristine image was noiseless, Gaussian noise was added to the real and imaginary parts of the k-space data used to generate the corrupted image. The scaling factor was chosen to be ε = 0.072, as was the case for the acquired data shown in Fig. 3a. The banding arising from the term linear in ε is apparent and only shows up where the intensity of the original image is high, in both measurement and calculation. The term quadratic in ε is more easily observed after the image window and level are adjusted to stress low signal regions, as shown in Fig. 3c,d. The quadratic term is generally too weak to be noticeable in images acquired in vivo. The observations are in agreement with Eq. [13].

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Figure 3. a: Image of fixed demarrowed trabecular bone specimen enclosed in a cylindrical container showing a pronounced banding artifact. b: Simulated image of a homogenous cylinder generated using Eq. [11] with ϵ = 0.072 calculated from the raw data of a. c,d: images a and b with the same display level and width adjusted to highlight the quadratic term under the square root of expression [13].

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We now show that two modifications of the original 3D FLASE sequence effectively eliminate the banding artifact. In the first modification all phase encodings are placed after the refocusing pulse, eliminating the phase encoding of the spurious magnetization that occurs before the first refocusing pulse (before t1 in Fig. 2). In this case the spurious magnetization is phase-encoded only once, after having been exposed to the second refocusing pulse (shaded region in Fig. 4a). Since both the main echo magnetization and artifact magnetization experience the same sequence applied gradients, the resulting gradient echoes are simultaneous. What would otherwise be an artifact-producing signal now adds a properly phase-encoded gradient echo to the main signal.

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Figure 4. a: Modification of the 3D FLASE sequence in which all phase encodings and the readout prephasing are performed after the refocusing pulse (shaded area). Since the magnetization resulting from the imperfect refocusing pulse is longitudinal while the phase encoding and prephasing gradients are played out, no k-space shift of the resulting echo occurs. b: Modification of the 3D FLASE sequence in which the sign of the crusher is matched to the sign of the slice-encoding gradient. The artifactual echo is thus always in the half of k-space opposite of the one in which the currently phase encoded signal lies.

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While the signal in this case is a linear combination of a spin and a gradient echo, the resulting adverse effect by introducing susceptibility dephasing is negligible, as will become clear in the following. The magnitude, ε, of the artifactual signal relative to the main signal in clinical imaging is on the order of 5%. Based on an analysis similar to the one that led to Eq. [13] in the previous section, it follows that the gradient echo will produce two terms corrupting the spin echo image

  • equation image(14)

where ρSE and ρGE are the images from the spin and gradient echo, respectively. The term linear in ε is weighted by the spin echo signal, thus mitigating the susceptibility dephasing effects in the gradient echo. The term quadratic in ε is negligible because ε is already small. Even though it seems that by this argument the artifactual signal should never be important, one has to keep in mind the position of the artifactual echo in k-space. In the original sequence the artifactual signal is dominant and much larger than the main signal at k-space frequencies at which the artifactual echo appears. This is evident in the cosine modulation of the main signal in Eq. [13]. In the case of the modified sequence both echoes are centered at k = 0, where the main signal is dominant. Having all phase encodings after the refocusing pulse also allows the spins to reach true steady state, since all zeroth gradient moments between every two RF pulses are now the same for every repetition period of the acquisition.

Although this situation is preferable, in practice a compromise has to be sought, since playing out all phase encoding after the refocusing pulse would prohibitively increase the echo time. Since we acquire a fractional echo in which the readout dephasing is relatively short, placing only that gradient after the refocusing pulse does not adversely affect the echo time. This modification thus differs from the one shown in Fig. 4a in that only the readout prephasing is placed after the refocusing pulse (shaded region of Fig. 4a). As in the original FLASE 3D (Fig. 1b), the z phase encoding is played out before the refocusing pulse and the y phase encoding is split. In this variant of the sequence the undesired magnetization is still phase-encoded twice in both y and z, as in the original 3D FLASE, but it is dephased only once in x, as in the ideal case described in the previous paragraph. The center of the artifactual echo is moved to kx = 0, while its features in ky and kz remain the same. This approach completely removes the banding in the readout direction (removes kprex terms in Eq. [11]) but still leaves some banding along the slice direction (Δkzz terms in Eq. [11]). However, the artifactual signal in this case is weakened for the following reason. As mentioned, the gradient echo of the artifactual signal occurs at the same time as the main echo, i.e., TE/2 after the refocusing pulse. Since this magnetization follows coherence pathway I in Fig. 1, the spin echo for the artifact signal still occurs a TE after the refocusing pulse, which is TE/2 later than the gradient echo. The gradient echo is thus weakened by a nonrephased T2′ dephasing present during the TE/2 time difference between the spin and gradient echo. Figure 5 shows a comparison between a scan with the original 3D FLASE and the modified sequence. The banding is clearly eliminated in the latter.

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Figure 5. a: 3D FLASE image of the wrist with the parameters given in Fig. 1, acquired using the original pulse sequence with the banding artifact present. b: Corresponding image acquired in the same subject using the modification of Fig. 4a. Both scans were performed with the same flip angle calibration; thus, the imperfection of the refocusing pulse was the same for both scans.

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For the image shown in Fig. 5a the relative magnitude of the artifactual echo was ε = 2%, while for that in Fig. 5b it was ε = 1.75%. Figure 6 shows the position of the artifactual echo in the case of the original sequence (Fig. 6a) and in the case of the above modification (Fig. 6b). While in Fig. 6b) the echo is not completely suppressed, it is diminished by 25% and moved to a different position in k-space (along the kz axis) according to our prediction. This is also one way to confirm that the source of the artifact is coherence pathway II. The in-plane banding is completely removed while the banding along the slice direction could only be observed in regions of low signal.

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Figure 6. a: A view of the raw data in the kxkz plane at ky = 0 for the scan shown in Fig. 5a. The artifactual echo is displaced from the center of k-space in both the slice and readout directions. b: The same slice of raw data for the scan shown in Fig. 5b. The artifact is now displaced only along the slice directions.

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A second possibility, illustrated by Fig. 4b, is to use the fact that the crusher gradient displaces the artifactual echo by Δkz along the kz direction in k-space. Let us assume that the sign of crusher gradients is the same as the sign of the kz phase encoding during a given repetition period. This would result in the artifact being phase-encoded always in the opposite half space along kz as the sequence is played out; thus, the artifact signal would effectively be dephased. Figure 7 shows how the artifactual echo is sampled when the corresponding half-spaces of the main signal are sampled. A 2Δkz-thick strip containing the strongest artifactual signal (k = 0) is never sampled for the artifactual echo. However, this method has the unwanted consequence of disrupting the steady-state magnetization even further, since the crusher gradients that produce part of the phase encoding in the nonsteady-state signal are now varied during the acquisition. The variation can be large, since the phase imparted by the crushers changes sign when kz changes sign. To minimize the effect of such a disruption of the steady state, the phase encoding along ky should be varied first and then the phase encoding along kz. This order of phase encodings leads to only a single change of the sign of the crusher gradients. In the opposite case, when kz is varied first, the crusher would change sign twice for every ky value.

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Figure 7. Parts of k-space sampled by the main and artifactual echoes using the 3D FLASE modification shown in Fig. 5b. The hatchings correspond to two different signs of the crusher and slice-encoding gradients applied in the modified sequence. Note that the artifactual signal is never phase-encoded near the center of k-space that has the highest signal amplitude.

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Finally, since the k-space position of the artifactual signal is well defined, it has the potential of being removable by postprocessing. The advantage of this approach is that it is amenable to previously acquired data. Toward this goal, three different ways of removing the artifact were implemented, all modifying the k-space data within an ellipsoid concentric with the artifactual echo: 1) by zeroing the k-space signal within the ellipsoid; 2) by scaling down the k-space signal amplitude smoothly to a constant level within the ellipsoid; and 3) by scaling down the k-space signal amplitude smoothly to zero at the center of the ellipsoid. While all of these methods remove the banding artifact, they, of course, also remove wanted signal present in the altered region of k-space. Data processed with the artifact removed by postprocessing showed slightly lower reproducibility of calculated structural parameters than data in which neither prospective nor postprocessing correction of the artifact was performed. This result suggests that prospective correction of the artifact is preferable.

CONCLUSIONS

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES

Short-TR 3D spin-echo imaging is prone to artifacts from imperfections in the phase-reversal pulse that can give rise to stimulated echoes that occur in subsequent pulse sequence cycles. These echoes give rise to low-frequency modulations in image space. A k-space analysis elucidated the exact nature of this artifact, which was found to be in excellent agreement with the experimental data, and two alternative strategies for alleviating the problem are demonstrated.

Acknowledgements

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES

The authors thank Aranee Techawiboonwong for generous help in performing some of the experiments.

REFERENCES

  1. Top of page
  2. Abstract
  3. THEORY AND METHODS
  4. RESULTS AND DISCUSSION
  5. CONCLUSIONS
  6. Acknowledgements
  7. REFERENCES