| INTRODUCTION |
The association between chronic low back pain and morphologic, structural changes of the paravertebral muscles has been widely acknowledged (1, 2, 3, 4). Considerable research effort has been directed at investigating the fatty infiltration of the paravertebral muscle, as well as the reduction of muscle volume, as a clinical predictor of post-surgical outcome (5) and spinal stability (6). Additionally, fatty infiltration of the muscle can also be caused by chronic conditions such as age-related sarcopenia (7), which is the generalized gradual loss of skeletal muscle mass and strength; it is associated to functional limitation, physical disability, and muscle strength (8, 9). Thus, the accurate quantitative measurement of paravertebral muscle fat is an important issue. Magnetic resonance (MR) imaging, which has been investigated for its ability to quantify fat by measuring fat-signal fraction (FF) in various tissues (10, 11, 12, 13), is useful for this purpose.
Prior studies have demonstrated that MR imaging techniques involving the quantitative assessment of FF using Dixon technique (14, 15) with two-echo (16) and three-echo (10, 11, 17) in phantom and variable tissues (mostly in the liver) are accurate tools (18). Recently however, FF measurement using multi-echo Dixon techniques have been developed to further improve the accuracy of this approach (19, 20, 21, 22, 23). However, there is still no consensus on whether signal acquisition at more echo time points, with or without T2*-correction, would result in more accurate FF measurement of the skeletal muscle.
Therefore, the purpose of our study was to assess whether T2*-corrected multi-echo acquisition allows for a more accurate FF measurement of the paravertebral muscle in comparison with FFs derived from non-T2*-corrected two-echo Dixon or T2*-corrected three-echo Dixon, using FF measurements from single-voxel MR spectroscopy (MRS) as the reference standard.
| MATERIALS AND METHODS |
Patient Population
This prospective study was approved by our Institutional Review Board prior to patient recruitment. Written informed consent was obtained before the enrollment of each patient. Between June 2013 and October 2013, 60 consecutive patients (39 women, 21 men; mean age: 54.3 ± 19.1 years; age range: 20-92 years) were included in this study. Inclusion criteria were: 1) age greater than 18 years, 2) agreement to participate in the study, and 3) a history of chronic low back pain for more than 3 months. Patients who refused to enroll in the study, those who were claustrophobic, had a cardiac pacemaker, and/or a history of spinal surgery or metallic implant placement, were excluded.
MR Imaging Acquisition
All MR imaging procedures were performed with a 1.5T MR imaging scanner (MAGNETOM Avanto, Siemens Healthcare, Erlangen, Germany) using an integrated spine matrix coil.
All subjects underwent the standard clinical MR imaging protocol at our institution, including T1-weighted turbo-spine echo (TSE) imaging in the sagittal plane, and T2-weighted TSE imaging in the axial and sagittal planes. Imaging parameters for this routine protocol are summarized in Table 1. Next, three different MR imaging pulse sequences for FF quantification were performed using prototype multi-echo three-dimensional gradient echo sequences with inline reconstruction (VIBE-Dixon, work-in-progress #432 for two-echo and three-echo, work-in-progress #798B for six-echo, Siemens Healthcare, Erlangen, Germany). The pulse sequences sampled two, three, and six echoes, respectively. T2*-corrected FF calculation was performed for three- and six-echo data, whereas T2*-correction was not possible for two-echo VIBE-Dixon sequence inherently. For the three-echo acquisitions, T2* was estimated in a first step by a log-linear fit from the two in-phase echoes. The first opposed-phase and the first in-phase echoes were then T2*-corrected, prior to 2-point Dixon water/fat separation and FF calculation. For the six-echo acquisitions, a multi-step adaptive fitting approach was performed with the technique described in a previous study (24), using a single peak fat model. This was because we could not guarantee that a multiple peak fat model, which has previously been tested in the liver, would be appropriate for the paravertebral muscle.
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Images from the protocols were obtained at levels L4 through L5 of the spine, in the axial plane. All fat and water signal-only images with T2* (three- and six-echo) or without T2* (two-echo) images were consequentially reconstructed, thereby automatically generating axial FF mapping images from non-T2*-corrected two-echo VIBE-Dixon, T2*-corrected three-echo VIBE-Dixon, and T2*-corrected six-echo VIBE-Dixon sequences. FFs were calculated with the equation, fat signal / (water signal + fat-signal), using the signal contributions from water and fat obtained in FF mapping images. The imaging time was 51 seconds for non-T2*-corrected two-echo, 86 seconds for T2*-corrected three-echo, and 161 seconds for T2*-corrected six-echo VIBE-Dixon sequence. The imaging parameters of these three pulse sequences are summarized in Table 2.
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A flip angle of 5° was used to minimize T1-relaxation effects, which affects the error in fat quantification and is influenced by the difference in T1 values between fat and water when gradient-echo pulse sequences are used (25, 26). As the reference standard, T2-corrected single-voxel multi-echo 1H MRS (HISTO; High speed T2-corrected multiple echo 1H MRS-Fat and R2 Quantification, work-in-progress #599B, Siemens Healthcare) (9) was performed using stimulated echo acquisition mode, with the following parameters: repetition time = 3000 ms, echo time (TE) = 12, 24, 36, 48, and 72 ms, 1024 acquired points, bandwidth = 1200 Hz, voxel size = 15 × 15 × 15 mm3. The acquisition duration for the HISTO sequence was 15 seconds.
A spectroscopic voxel was placed in either the lumbar erectus spinae muscle or lumbar multifidus muscle, at any random level between the L4 and L5 vertebral bodies, and repeated in the contralateral side in every subject. The voxel was placed by an MR imaging technician under the supervision of a musculoskeletal radiologist, who had around 30 years of experience in the field of musculoskeletal radiology. We focused on the lumbar erectus spinae muscles and lumbar multifidus muscles, which are known to be important stabilizers of the lumbar spine (27) at levels L4 and L5; this selection was based on an earlier study that demonstrated that the paravertebral muscle fat at the L4 through L5 levels could be a useful marker of whole body skeletal muscle fat (28). Based on the signals from fat and water at each TE, FF was automatically calculated from this sequence. An example of the spectral peaks at a certain TE and the automatically calculated FF from the HISTO sequence is demonstrated in Figure 1.
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Image Analysis
Fat-signal fraction from the mapping images were measured by drawing a circular region of interest (ROI) at the MR imaging workstation (VB17, Siemens Healthcare) using Syngo software. After randomization of the subjects, two musculoskeletal radiologists (observer A, observer B) performed all measurements independently, without knowledge of the FF measurements from the MRS. ROIs were drawn on a mapping image from the two-echo sequence, at the same location of the spectroscopic voxel that referred to the captured images obtained while placing the voxel in each instance of MR imaging. These were then copied and pasted to two other mapping images (Fig. 2). The diameter of the circular ROI was fixed at 15 mm, based on the voxel size of MRS. For precise representative data acquisition over the entire voxel area, ROIs were drawn at the three consecutive slices of mapping images in the same manner as stated above. As a result, three ROI measurements from the three FF pulse sequences were obtained for every voxel location.
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Statistical Analysis
All continuous data were expressed as mean ± standard deviation values. The intraclass correlation coefficient (ICC), with a two-way random model of absolute agreement, was used to evaluate the agreement of FF measurements between the two radiologists (observer A and B). For interpretations of the results, ICC values less than 0.40 indicate poor reproducibility, ICC values in the range of 0.40 to 0.75 indicate fair to good reproducibility, and ICC values greater than 0.75 indicate excellent reproducibility (29).
To determine the agreement between the FF measurements obtained from the FF maps from VIBE-Dixon sequences and those obtained using MRS, Lin's concordance correlation coefficient (CCC) was used (30). Accordingly, three CCCs for the VIBE-Dixon sequences were obtained for measurements by each observer. Subsequently, multiple comparisons with Streiger's Z-test was done as post-hoc analysis through a Bonferroni correction by using the aforementioned CCCs, to determine if there was a significant difference between the results obtained from the three different VIBE-Dixon MR sequences, with an adjusted significance level (0.05/3).
In order to show the mean measurement bias (mean difference) with corresponding limits of agreement of the FF measurements obtained from the three different VIBE-Dixon sequences and those from MRS, Bland-Altman analysis was used with a 95% confidential interval (31). All statistical analyses were performed using statistical software (SAS Institute, version 9.2, Cary, NC, USA).
| RESULTS |
A total of 120 spectroscopic FF measurements were done. Of these, 117 spectroscopic FF measurements were completed either bilaterally (57 of 60) or unilaterally (3 of 60), and the values were used as reference standards. Three MRS data from three subjects were missed owing to a storage error while saving the result; hence, the measurements from the remaining 117 sites were assessed.
The mean spectroscopic FF percentage was 14.3 ± 11.7% (range, 1.9-63.7%). The mean FFs from each VIBE-Dixon sequence obtained by observers A and B are summarized in Table 3. The interobserver agreement for the FF measurements between observers A and B was excellent for all three VIBE-Dixon sequences (Table 4).
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Lin's CCCs between the spectroscopic and all the imaging-based FFs were statistically significant (p <0.001) (Fig. 3). The CCCs for each VIBE-Dixon sequence in observer A and B are summarized in Table 4.
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As compared with the other imaging-based FFs through multiple comparisons, the FF obtained from T2*-corrected six-echo VIBE-Dixon sequence showed the best agreement with the spectroscopic data, and was statistically significant (p <0.001). However, there was no statistically significant difference between non-T2*-corrected two-echo VIBE-Dixon and T2*-corrected three-echo VIBE-Dixon in observer A (p = 0.44) or observer B (p = 0.84).
Figure 4 shows the mean measurement bias with corresponding limits of agreement for FFs obtained from three different VIBE-Dixon sequences and those from MRS. The mean measurement bias was -2.5% (range, -9.2% to 4.2%) for two-echo, -2.4% (range, -8.6% to 3.7%) for three-echo, and 0.2% (range, -2.8% to 3.3%) for six-echo methods as evaluated by observer A, and -2.6% (range, -9.0% to 3.9%) for two-echo, -3.1% (range, -9.4% to 3.3%) for three-echo, and 0.5% (range, -2.9% to 3.9%) for sixecho methods as evaluated by observer B. In the findings of each observer, the range of limits of agreement was very narrow for T2*-corrected six-echo VIBE-Dixon findings, in comparison with the other sequences.
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| DISCUSSION |
We attempted to validate whether T2*-correction, based on signals from multiple TE, helps to increase the accuracy in the measurement of FF, especially for paravertebral skeletal muscle. Several researchers have focused on quantification of skeletal muscle fat by MR imaging in various parts of the body, including the rotator cuff (12, 32), thigh (25), and paravertebral muscles (33) using Dixon technique with or without T2*-correction. However, to the best of our knowledge, our research is the first one that demonstrates that FF obtained by T2*-corrected six-echo VIBE-Dixon best correlates with MR spectroscopic FF data, in comparison with inherently non-T2*-corrected two-echo VIBE-Dixon and T2*-corrected three-echo VIBE-Dixon data.
Most of the previous FF studies in the liver have already shown that T2*-correction generally improves the accuracy of FF measurement in three or more TEs using multi-echo VIBE-Dixon technique (34, 35, 36, 37). Previous studies on the liver have also established that three-echo VIBE-Dixon shows a more accurate measurement in the assessment of hepatic FF than two-echo VIBE-Dixon, and that six-echo VIBE-Dixon shows a more accurate measurement than three-echo VIBE-Dixon (34, 38). We have confirmed that the correlation between spectroscopic data and all imaging-based FFs was statistically significant, which was consistent with the results of a previous investigation, that assessed the FFs of lumbar paravertebral muscles from three different gradient echo sequences based on Dixon techniques using non-T2*-corrected two-echo, T2*-corrected multi-echo, and non-T2*-corrected multi-echo (33). However, Fischer et al. (33) unexpectedly found that T2*-correction was not helpful. They demonstrated that the T2*-corrected multi-echo sequence data did not correlate better with FF data from spectroscopy, than the non-T2*-corrected multi-echo sequence did; they hypothesized that this could be owing to the reduction of the signal-to-noise ratio caused by T2*-correction. Another report from Gaeta et al. (19) had explained that T2*-correction had no more benefit than non-T2*-corrected method in skeletal muscle than for tissues with a longer T2*, referring that the T2* of muscle is around 30 ms.
Although the necessity of T2*-correction in the assessment of skeletal muscle fat has not been established, we employed T2*-correction for three-echo and six-echo VIBE-Dixon sequences from the commencement of our study, based on the hypothesis that there are potential factors that can cause T2* decay in muscle. It is well known that the cause of T2* decay in hepatic parenchyma can mainly be explained by the incidence of iron (39). Theoretically, there are non-negligible amounts of iron stored in skeletal muscle that can affect the local magnetic inhomogeneity in the muscle, as is the case in the liver (40). Recent studies, however, have revealed that iron also plays an important role in muscle atrophy (41) and in the development of sarcopenia (42). This process can be more complex in the paravertebral muscles of patients with chronic low back pain when combined with microstructural changes of muscle fiber type (4) and changes of muscle fiber stiffness (43). We believe that such changes should not be neglected in the skeletal muscle, which may influence the local magnetic inhomogeneity, although the influence can be subtle with individual variances.
It is noteworthy that a prior phantom study assessed FF with fat-water-iron phantom using T2*-corrected multi-echo chemical shift-based method, with a variable iron concentration (37). This phantom study demonstrated the necessity of T2*-correction in six-echo chemical shift-based fat-water separation method, regardless of the iron concentration (0-50 µg/mL), particularly in conditions with FF lower than 50%. Since most of the observed FFs were lower than 50% in the present study, we assumed that T2*-correction could be helpful in the accurate measurement of muscle FF, even without knowing the exact iron concentration in muscle. Accordingly, our study result with T2*-corrected six-echo sequence has confirmed the prior phantom study result in vivo, particularly in the skeletal muscle.
In this study, the narrow range of limits of agreement in T2*-corrected six-echo VIBE-Dixon data was small enough to explain the more precise correlation with FF derived from MRS according to Bland-Altman analysis. Underestimation of FF measurement was found in all three imaging-based FFs with each observer, but was least prominent in T2*-corrected six-echo sequences. We observed slight errors in the FF measurements in each VIBE-Dixon sequence. However, most of the measurements were within the 95% limits without systematic error in the assessment of FF. For each observer, the smallest mean difference (0.2% in observer A, 0.5% in observer B) was noted in T2*-corrected six-echo VIBE-Dixon, which enhances confidence in this methodology.
We also observed that there was no statistically significant difference between the non-T2*-corrected two-echo VIBE-Dixon and T2*-corrected three-echo VIBE-Dixon sequences. This might have been caused by insufficient T2*-correction in three-echo VIBE-Dixon sequence. Another plausible explanation can be inferred from Reeder et al. (44), who suggested that it is necessary to acquire at least six echoes for the optimal separation of water and fat signals with T2*-correction (20).
Our study has several limitations. First, we used MRS as a reference standard rather than traditional pathologic confirmation by muscle biopsy. However, muscle biopsies cannot always be performed, owing to ethical problems. Additionally, the small portion of muscle gained from a biopsy cannot always represent the whole muscle condition. Secondly, there was a relatively small population of patients with a higher muscle FF (> 50%). Therefore, we still need to elucidate whether this methodology can be applied for patients with massive fat infiltration. Third, our study used single fat-peak spectral modeling for multi-echo VIBEDixon MR imaging. It has been suggested by Yokoo et al. (34) that a single fat-peak model could be inadequate for reliable T2* estimation in MRI. However, little has been known about the usefulness of multi-peak fat modeling for muscle FF measurement in comparison with single-peak fat modeling. Therefore, further large-scale studies would be needed, taking into consideration both T2*-correction and multi-peak fat modeling.
In conclusion, T2*-corrected six-echo VIBE-Dixon sequence agrees with spectroscopic FFs better, as compared to non-T2*-corrected two-echo and T2*-corrected three-echo VIBE-Dixon sequences. Based on our observations, T2*-corrected six-echo VIBE-Dixon sequence could be used as an accurate, clinically feasible, noninvasive quantification tool for muscle fat quantification in lumbar spine MR imaging.
Notes
This research was supported by the Basic Science Research Program through the National Research Foundation (NRF) of Korea funded by the Ministry of Education, Science and Technology (NRF-2012R1A1A2007991).
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Yeon Hwa Yoo
Department of Radiology, Gangnam Severance Hospital, Yonsei University College of Medicine, Seoul 06273, Korea
Hak-Sun Kim
Department of Orthopedic Surgery, Gangnam Severance Hospital, Yonsei University College of Medicine, Seoul 06273, Korea
Young Han Lee
Department of Radiology, Severance Hospital, Yonsei University College of Medicine, Seoul 03722, Korea
Choon-Sik Yoon
Department of Radiology, Gangnam Severance Hospital, Yonsei University College of Medicine, Seoul 06273, Korea
Mun Young Paek
Healthcare Sector, Siemens Ltd., Seoul 03737, Korea
Hanna Yoo
Biostatistics Collaboration Lab, Yonsei University College of Medicine, Seoul 03722, Korea
Stephan Kannengiesser
Healthcare Sector, Siemens AG, Erlangen 91052, Germany
Tae-Sub Chung
Department of Radiology, Gangnam Severance Hospital, Yonsei University College of Medicine, Seoul 06273, Korea
Ho-Taek Song
Department of Radiology, Severance Hospital, Yonsei University College of Medicine, Seoul 03722, Korea
Jin-Suck Suh
Department of Radiology, Severance Hospital, Yonsei University College of Medicine, Seoul 03722, Korea
Sungjun Kim
Department of Radiology, Gangnam Severance Hospital, Yonsei University College of Medicine, Seoul 06273, Korea
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