With increasing utilization and improved resolution of computed tomography (CT), an increasing number of pulmonary part-solid nodules (PSNs) of uncertain malignant potential are being detected.1–8 CT imaging characteristics to reliably discriminate benign from malignant lesions are lacking.3,4,8–21 Whether the presence of a solid component is suggestive of malignancy, and its prognostic implications, remains controversial.8,19–21 Studies such as those by Oh et al17 and Lee et al20 reported that as many as 69.8% of PSNs are transient. Positron emission tomography (PET) with F-18 fluorodeoxyglucose (18F-FDG) has not been proven to be as valuable for evaluating PSNs.22,23 Malignant subsolid nodules can be falsely negative on FDG-PET.22 Despite these limitations, patients often undergo serial follow-up CT or PET/CT examinations and/or biopsies or wedge resections because of the malignant potential of PSNs. Disadvantages of this approach include invasive procedures, postprocedural morbidity, and significant additive radiation.24–26
Magnetic resonance imaging (MRI) has been explored as an attractive alternative to CT for pulmonary nodule evaluation because of the lack of ionizing radiation.26–32 T1-weighted and T2-weighted imaging has been demonstrated to differentiate PSNs from normal lung.33–36 However, the role of diffusion-weighted imaging (DWI) for nodule evaluation is still unclear, with conflicting evidence from several studies.34–36 These results have yet to be verified by other studies. The majority of prior studies employing DWI utilized a 1.5 T system, and results might be different with a 3 T system. Besides DWI, T2* values, which are sensitive to microscopic magnetic field inhomogeneities, might also be related to lesion density but have not yet been applied to nodule characterization. Studies to date have encompassed pulmonary nodules ranging from pure ground-glass opacities to solid lesions, with most targeting solid nodules. None have focused specifically on PSNs.
The purpose of our prospective pilot study was to assess whether MRI characteristics such as T1-weighted and T2-weighted signal intensity, apparent diffusion coefficient (ADC), and T2* values obtained with a 3 T system can distinguish benign from malignant PSNs and predict the aggressiveness of the latter. In addition, we sought to compare these MRI parameters with retrospectively retrieved morphologic and physiological values derived from CT and PET/CT. We hypothesize that these MRI parameters can discern benign from malignant PSNs and predict the aggressiveness of the later while being comparable to conventional examinations such as CT and PET/CT.
MATERIALS AND METHODS
This institutional review board–approved prospective study conformed to Health Insurance Portability and Accountability Act guidelines on patient data integrity. Written informed consent was obtained from each of the consecutively recruited patients. Inclusion criteria consisted of age 18 years or older and presence of 1 or more PSNs>1 cm in diameter as seen on CT and/or PET/CT scheduled for surgical resection or biopsy between July 2011 and June 2014. Exclusion criteria consisted of claustrophobia and non–MRI-compatible implanted metal objects. Thirty-three patients (23 women) with ages ranging from 33 to 96 years (mean age 73.5±13.8 y) were recruited (Table 1). Data on pathology were unavailable for 5 patients before final analysis, and these cases were excluded from evaluation, yielding a total of 28 patients with 32 PSNs for final analysis (4 patients had >1 nodule). Eighteen lesions in 17 patients were malignant. All malignant lesions were adenocarcinoma, except for 1 squamous cell carcinoma (Table 1). Seven malignant lesions were acinar, and 9 were lepidic subtypes. Ninety-four percent (15 of 16) of malignant lesions were poorly differentiated. All pathology specimens obtained at the time of biopsy or resection were processed into paraffin-embedded sections and assessed as per routine surgical pathology practice.
Unenhanced pulmonary MRI examinations were carried out using 3.0 T whole-body scanners (Siemens Tim Trio and Siemens Magnetom Verio, Siemens Medical Solutions, Erlangen, Germany) with a maximal imaging field of view of 50 cm. Patients were imaged with either a 32-channel torso array coil (Rapid Biomedical, Rimpar, Germany) with 16 elements, each situated anteriorly and posteriorly, or an 18-channel system with a combination of 6-element anterior body matrix coil and 12-element posterior spine matrix coil. Parallel imaging was utilized for both T2-weighted and DWI with an acceleration factor of 2. T2-weighted anatomic localization images of the chest were obtained during free breathing in the axial, coronal, and sagittal planes (Table 2), followed by targeted axial imaging of the lesion at end inspiration with breath-holding utilizing T1-weighted 2-dimensional (2D) gradient echo and T2-weighted 2D turbo spin echo (TSE) with fat saturation sequences. Combined acquisition time for the aforementioned sequences was approximately 5 minutes. Subsequently, T2*-weighted multiple gradient echo axial images of the target lesion were obtained with multiple echoes at end inspiration with acquisition time of approximately 5 minutes. Finally, diffusion-weighted 2D TSE axial images were obtained at end inspiration with breath-holding using a single-shot, twice-refocused spin echo sequence with bipolar diffusion gradients and centric-ordered TSE readout. Although most previous lung DWI studies have used echo-planar imaging (EPI) readouts, the TSE sequence was chosen to avoid susceptibility-induced distortions that worsen at higher applied fields and are particularly pronounced in the lung. Diffusion sensitization was in the anterior-posterior direction applied with weighting factors of b=0, 50, 100, 150, 200, 300, 400, 500, and 600 s/mm2. Acquisition time for these sequences was 5 to 10 minutes. Images of a PSN in various MRI sequences are illustrated in Figure 1. Utilizing the localization MRIs and the previously obtained CT or PET/CT as guides, we typically obtained about 1 to 3 slices per sequence that encompassed the lesion. If >1 slice was obtained, the slice that best represented the nodule (typically one with the best signal to noise ratio) was utilized for analysis.
Diagnostic unenhanced clinical thoracic CT examinations were performed utilizing 64-slice or 128-slice multidetector computed tomography scanners (Definition, Definition AS+ or Definition Flash, Siemens AS, Malvern, PA) with protocols using 120 to 140 kVp, gantry rotation time of 0.33 to 0.5 seconds, 512×512 matrix, and reference tube current of 70 to 160 mAs. The images were reconstructed with either b46 or b60 kernels and a section thickness ranging from 1 to 1.5 mm at 0.8 to 1 mm intervals.
Patients were instructed to fast starting midnight before intravenous administration of 15±1.5 mCi of 18F-FDG. PET imaging was performed 60 to 70 minutes after FDG administration with integrated PET and 64-slice or 16-slice CT devices (Discovery 690 or Discovery RX; GE Medical Systems). All images were acquired during shallow breathing. The acquisition time for PET in 3D and time-of-flight mode was 3 to 5 minutes per table position. PET data were resampled from either a 128×128 or a 192×192 matrix to a 512×512 matrix to match the CT data to allow attenuation correction, and a CT transmission map was generated. PET image data were reconstructed iteratively using the ordered subsets expectation maximization algorithm with segmented attenuation correction (2 iterations, 35 to 36 subsets) and the CT data. The CT protocol used 120 kVp, gantry rotation time of 0.5 seconds, a pitch of 1 (1.5 for the 16-slice CT), and adjustable reference tube current depending on the patient’s body mass index, ranging from 20 to 90 mAs (maximum 120 mAs for the 16-slice CT) for a standard patient. A de-noising algorithm was applied to CT images of patients with body mass index >35. The transaxial CT slice thickness was 3.75 mm with an interval of 3.27 mm for fusion with the transaxial PET images. The PET, CT, and fused images were available for review in the axial, coronal, and sagittal planes by means of software (Advantage Windows; GE Medical Systems) on a computer workstation.
MR examinations were uploaded onto a stand-alone workstation (Leonardo VE30A, Siemens Healthcare) for analysis. Regions of interest (ROIs) were independently placed by a second-year medical student (V.T.) after a training session of 5 cases and by 1 thoracic radiologist with 7 years of thoracic imaging experience (C.W.K.), both blinded to clinical data at the time of evaluation. The ROIs encompassed the entire lesion with the highest signal intensity for the T1, T2, T2*, and diffusion-weighted sequences (Fig. 2A). ROI size was kept constant over the range of b values for the DWIs. Signal intensities derived from the ROIs as a function of b value were fitted to a monoexponential decay model whose decay rate was the ADC as derived from the following equation, with ADC and S0 as free parameters: Sb=S0 exp (−b×ADC), where Sb is the signal intensity for a given b factor, and S0 is the signal intensity without diffusion weighting. For T2* values, ROIs were drawn over each lesion to determine the signal intensity for each TE. Signal intensity at each echo time, SE, is given as follows: SE=S0 exp (−T2*/TE), where S0 is the signal at zero TE. T2* value was calculated by nonlinear least-squares fitting of the above equation to measured values of SE, with T2* and S0 as free parameters. ROIs were redrawn at least 2 weeks later by each reader and the analysis repeated.
A single reader (C.W.K.) performed all CT assessments. For CT Hounsfield unit (HU) measurements, ROIs were placed once to encompass the entire lesion (Fig. 2B). Similarly, an ROI was drawn to encompass the solid component, which was defined as the part of the lesion where the normal pulmonary architectures (vessels, airways, and alveoli) were completely obscured. The maximal diameter on axial images in which the nodule appeared the largest was determined. The longest linear dimension of the solid component was also recorded on the axial section in which the solid component appeared the largest. Percentage solid component was visually stratified into 4 groups: <25%, 25% to 49%, 50% to 75%, and >75%.
Statistical analysis, detailed below, was performed to assess the relationship between ADC, T2*, and measured signal intensity values on T1-weighted and T2-weighted images and benignity or malignancy. For patients with pathologically proven adenocarcinoma, MRI parameters were correlated to degree of invasiveness and histologic subtype. Statistical correlation between the ADC, T2*, T1, and T2 values and maximum standard uptake values (SUVmax) was performed in patients with preoperative PET or PET/CT examinations. MRI parameters were assessed for correlation with HU measurements as well. Lesion visibility, defined as whether or not a lesion was seen on a particular MRI sequence, was also correlated with solid component size and percentage of solid component on CT.
Statistical analyses were performed with the SAS System (version 9.4; Cary, NC). The analysis was conducted on a per-nodule basis under the assumption of statistical independence, accounting for a limited number of nodules nested within cases (32 nodules in 28 cases). Given multiple measurements that resulted from 2 measurements by each of 2 readers, the mean of the 4 values was utilized for diagnostic performance and correlation analysis. Continuous variables were reported as mean (SD) and categorical variables as frequency (percentage). The Spearman rank order correlation coefficient was utilized to assess for degree of correlation. Receiver operating characteristic curve analysis was performed to predict the association between malignancy and MRI parameters and to obtain sensitivity, specificity, and accuracy measurements. Intraclass correlations (ICC) were calculated to validate the intrareader and interreader reliabilities between readings for each rater. An overall ICC was also computed to summarize the variability of the 4 measurements within each nodule using a random-effects model with a random nodule effect. As proposed by Landis and Koch,37 ICC values >0.8 were regarded as “almost perfect,” those between 0.61 and 0.8 as “substantial,” those between 0.41 and 0.6 as “moderate,” and those between 0.21 and 0.4 as “fair.” Statistical tests with a 2-sided P≤0.05 were considered significant.
The mean ADC for benign nodules ranged from 1.1 to 1.3 μm2/ms, whereas malignant lesions ranged from 1.6 to 2.3 μm2/ms (Table 3). Mean T2* values for benign nodules ranged from 13.2 to 16.4 ms, and those for malignant nodules ranged from 12.2 to 24.2 ms. Mean T1 signal intensities for benign nodules ranged from 75.9 to 87.5, and those for malignant lesions ranged from 78.1 to 87.6. Mean T2 signal intensities for benign lesions ranged from 91.9 to 104.5, and those for malignant lesions ranged from 99.6 to 110. Of the 32 nodules, 32 were visualized on T1-weighted, 31 on T2-weighted, 30 on DWI, and 29 on T2*-weighted sequences.
The mean CT density for benign lesions was −396.6 HU, and that for malignant lesions was −424.0 HU. The mean CT density of the solid component for benign nodules was −226.9 HU, and that for malignant nodules was −248.6 HU. Data for solid component on CT were available for 30 lesions, as 2 outside CT results were unavailable at the time of data analysis. Four lesions contained <25% solid components, 7 contained 25% to 49% solid components, 9 contained 50% to 75% solid components, and 10 had >75% solid components. The mean diameter (SD) of the solid component for benign nodules was 1.3 cm (0.8), and that for malignant nodules was 1.5 cm (0.9). The maximum standardized uptake values were available for 4 benign lesions with a mean (SD) of 3.0 (2.3) and for 9 malignant lesions with a mean of 2.8 (1.7).
ADC significantly correlated with malignancy (P=0.005) with high discrimination (area under the curve 0.79; 95% confidence interval [CI], 0.63-0.96). An ADC≥1.28 μm2/ms predicted malignity with a pooled sensitivity of 83.3% (95% CI, 60.8%-94.2%), specificity of 66.7% (95% CI: 39.1%-86.2%), and accuracy of 76.7% (95% CI, 59.1%-88.2%). Neither T2*, T1, and T2 values, SUVmax, HU of the solid components, nor HU of the entire nodule correlated with malignancy (P>0.28). For adenocarcinomas, ADC inversely correlated with lepidic subtype (P<0.03) but not with acinar predominance (P>0.71, Table 4). Conversely, T2* correlated with the acinar (P=0.013) but not with the lepidic subtype (P=0.71). T1 and T2 did not correlate with histopathology subtypes. None of the MRI parameters significantly predicted tumor differentiation (P>0.11). For patients who had undergone PET imaging, we did not find significant correlation between SUVmax and any of the MRI parameters (P>0.56). T1 values showed significant correlation with HU measurements of the entire nodule (P<0.001) and with HU measurements of solid components (P=0.031), whereas ADC and T2* did not. T2 values showed significant correlation with HU of the entire nodule (P=0.024) and with HU of the solid components (P=0.008). The greatest diameter of the solid component correlated with lesion visibility on DWI and T2* images (P<0.04, Table 4). There was a significant correlation between percentage of solid component and lesion visibility on T1 (P=0.018).
Reader 1 demonstrated almost perfect intraobserver correlation for T1 values, substantial correlation for T2 and T2* values, and moderate correlation for ADC (Table 5). There was near-perfect intraobserver correlation for ADC, T1, and T2 values and moderate correlation for T2* evaluations for reader 2. We demonstrated almost perfect interobserver correlation for T1 and T2 values, substantial correlation for ADC, and fair correlation for T2* evaluations.
In summary, we have demonstrated that ADC significantly correlated with malignancy with high discrimination. ADC and T2* correlated with adenocarcinoma subtypes. No MRI parameters predicted tumor differentiation. SUVmax did not correlate with any MRI parameters. Visibility on T1-weighted images correlated with percentage of solid components. T1 and T2 values showed significant correlation with HU measurements of the entire nodule and with HU measurements of solid components.
Our results confirmed prior reports promoting MRI as a useful noninvasive means for imaging pulmonary nodules without the use of ionizing radiation.27–42 Although these studies used DWI with a 1.5 T system, our study represents one of the few studies utilizing a 3.0 T system. According to Ohba et al,39 both 1.5 and 3.0 T DWI modalities are equally effective for assessing pulmonary nodules. We also expected the higher field system provided by a 3.0 T system to generate a better signal to noise ratio and faster scan times compared with a 1.5 T system. The increased signal to noise ratio was helpful for the low-intensity PSNs, whereas the faster scanning time enabled easier breath-holding for the patients with pulmonary disease. Our study utilized a single-shot, twice-refocused spin echo sequence with bipolar diffusion gradients and centric-ordered TSE instead of EPI readouts to avoid susceptibility-induced distortions that worsen at higher applied fields and are especially pronounced in the lungs. According to Zhou et al,43 EPI is sensitive to main magnetic field heterogeneity, especially in the phase encoding direction, because of relatively low bandwidth. The use of spin echoes in the TSE sequence rather than gradient echoes in an EPI readout alleviates such an issue given the lessened phase encoding.
Although the numerical value of the ADC threshold demonstrated in this study is similar to previously reported thresholds,34–36,39,40,42 unlike most previous studies we found the mean ADC of malignant lesions to be higher than that of benign lesions. According to Liu et al,40 the mean ADC of malignant pulmonary lesions was significantly lower than that of benign pulmonary lesions in a cohort of solid nodules. This might be explained by the fact that our benign lesions tended to be very cellular, consisting of lymphoid hyperplasic, fibrotic, granulomatous, and other inflammatory nodules. In contrast, our malignant lesions consisted of PSNs that are not as cellular as the solid nodules discussed in prior studies. This is also supported by the fact that the mean HU of benign lesions was greater than that of malignant ones, indicating that the benign lesions were denser than the malignant ones, hence with increased barriers to water diffusion. In addition, most of our malignant lesions were high-grade adenocarcinomas, which might have necrotic components. Herneth et al44 reported that the ADC values of necrotic tumors are significantly higher than that of hypercellular neoplasia and benign lesions in mice. Similar to prior studies, our ADC threshold also demonstrated a high sensitivity of 88.9% with good accuracy (area under the curve 0.79), and the ADC correlated with adenocarcinoma histopathology subtype.41
To our knowledge this is the first study to examine the relationship between T2* and pulmonary nodule malignant potential. As expected, the mean T2* values and T2 signal intensities are higher for malignant lesions. This could be due to less susceptibility and increased water content in the malignant lesions, respectively. Our study showed that T2* correlated with acinar subtype of adenocarcinomas, which might be related to increased susceptibility and field inhomogeneity associated with this gland-forming adenocarcinoma subtype. As different adenocarcinoma subtypes have different prognostic significance levels,9 such a finding is clinically significant. We were unable to show that MRI parameters significantly predicted aggressiveness, but that could be because nearly all (15 of 16) malignant lesions were poorly differentiated.
To our knowledge, no study to date has associated T1 and T2 signal intensities and density measurements of PSNs. Given that nodule visibility is likely related to lesion cellularity, it is not surprising that the T1 and T2 signal intensities showed significant correlation with CT density of the entire PSN and CT density of solid components. Intuitively, one might think that solid component size might be related to lesion density; thus, visibility of these lesions on MRI sequences, such as on T1-weighted and T2-weighted images, might correlate with the solid component diameter; however, this was not the case in our study. This might simply indicate that nodule densities are more complex. However, nodule visibility, at least on T1-weighted images, correlated well with the percentage of solid components.
De Hoop et al18 reported that mass measurements allowed detection of nodule growth earlier than diameter, area, or volume measurements. HU was utilized in a formula to calculate mass. However, no study to date has directly correlated HU with malignancy, and we found that there is no direct association between HU and malignant potential. As expected, the greatest diameter of the solid component correlated with lesion visibility on DWI and T2* images.
Interestingly, the widely utilized SUVmax also did not correlate with malignancy. Although this could be due to the small sample size, inherent properties of PSNs differing from solid nodules is the more likely explanation, as Hattori et al45 demonstrated that PSNs had a mean SUVmax that was lower than the normally accepted SUVmax. None of the MRI parameters correlated with SUVmax, which is likely related to increased susceptibility effects, such as image artifacts and shortened T2* values, with a 3 T system, as discussed in the study by Ohba et al.39
Our study showed moderate to near-perfect intraobserver agreements for all MRI parameters and substantial to near-perfect interobserver agreement for ADC, T1, and T2 values only, which might be due to difficulty in visualizing these nodules on T2* images. The fact that T2* dephasing is more prominent at 3 T may have also contributed to technical difficulty in imaging these PSNs.
Our study had several limitations. Given that this is a pilot study, the small sample size limited our ability to demonstrate statistical significance for several relationships and precluded statistical adjustment for nesting of nodules within cases. As the sample size was considered too small to model the ROC using a multirater, multicase design, we elected to simplify the data by means of computing a composite score. Further study with a larger cohort size will eliminate these limitations. Second, respiratory and electrocardiography triggering were not utilized, which might have improved our image quality. However, we felt that the quality of breath-holding for the most important sequences was adequate, and the majority of our lesions were not adjacent to the heart or great vessels. Two of the 32 PSNs were not visualized on the DWI. One of these lesions was an elastic fibrous scar, and the other lesion was an aggregate of alveolar cells and macrophages, which likely lacked significant freedom for water diffusion to generate signals. Three of the 32 lesions were visualized on the T2* sequences by only 1 of the 2 readers. Two of the lesions were low-grade adenocarcinomas with lepidic growth, and the other nodule was an adenocarcinoma with significant pseudocavitation that might not have filled the air spaces with enough cells to reduce the field inhomogeneity enough for significant visibility on T2* imaging. Finally, we used b values for ADC measurements where the perfusion fraction was not excluded, and because only monoexponential fitting was performed the ADC values might have been slightly overestimated; however, the reproducibility should have remained the same.
In conclusion, 3 T MRI shows promise for PSN evaluation with ADC, providing a potential quantitative means for distinguishing benign from malignant lesions. ADC and T2* values can potentially be helpful for predicting adenocarcinoma subtypes. T1 and T2 values showed significant correlation with HU measurements of the entire nodule and of the solid components.
The authors acknowledge the following individuals for their contributions to this study: Christine U. Lee, MD, PhD, and David L. Levin, MD, PhD, for general oversight of the study; and Kimberly K. Amrami, MD, for administrative support.
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