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. Author manuscript; available in PMC: 2022 Jun 4.
Published in final edited form as: J Bone Joint Surg Am. 2021 May 19;103(10):887–899. doi: 10.2106/JBJS.19.00419

Large lytic defects produce kinematic instability and loss of compressive strength in human spines: an in-vitro study

Ron N Alkalay 1, Robert Adamson 1, Alexander Miropolsky 2, Roger B Davis 1, Mike L Groff 3, David B Hackney 1
PMCID: PMC9167060  NIHMSID: NIHMS1810659  PMID: 33755638

Abstract

Background:

In patients with spinal metastases, kinematic instability is postulated as a predictor of pathologic vertebral fractures. However, the relationship between this kinematic instability and the loss of spinal strength remains unknown.

Methods:

Twenty-four 3-level thoracic and lumbar segments from eight female cadaver spines, age 47–69 years, were kinematically assessed in axial compression (180N) and axial compression with flexion or extension moment (7.5Nm). Two patterns of lytic defects were mechanically stimulated; a) vertebral body defect (Taneichi1 model C, n=13) and b) model C + compromise of the ipsilateral pedicle and facet joint (Taneichi1 model E, n=11). The kinematic response was re-tested, and compression strength was measured. We applied two-way repeated measures analysis to test the model’s (C, E) effect on the change (lesion vs. control) in the segment’s kinematic response. We used multivariable linear regression to test the association between changes in the segment’s kinematic parameters and compressive strength.

Results:

Compared to model C, model E caused significantly higher changes to the segments extension (p=0.0233) and torsional (p=0.0429) ROM, and sagittal translation (p<0.001) under extension moments and sagittal translations (p=0.0452) under flexion moments. Destruction of the middle column (p<0.0001), higher axial translations under flexion (p=0.0002), and extension (p=0.0021), and sagittal translations under flexion (p=0.0009), were negatively associated with lower segment’s compressive strength.

Conclusions:

Critical spinal lytic defects affect kinematic abnormality with lower spine compressive strength, suggesting reduced spinal structural rigidity.

Keywords: Lytic defects, Kinematics, Mechanical test, Spinal instability

INTRODUCTION

The spinal column is the most common site of osseous metastases2, with up to 73% of breast and 68% of prostate cancer patients found to have spinal metastasis at autopsy3, 4. Bone metastases5 disrupt vertebral strength and rigidity6–8. This effect has its most dramatic manifestation when it causes pathologic vertebral fractures (PVF)9, affecting up to 30% of vertebrae with bone metastases10. Both the initial fracture and any subsequent collapse can lead to severe consequences, with up to 50% of patients with PVF suffering neurological deficits with further complications that may be fatal10. Currently, the treatment of these fractures is primarily reactive after a fracture has occurred. Enhancing the ability to predict PVF before they occur is critical, and as of yet, imprecise11, 12, element in the management of these patients13.

The impact of spinal metastasis on the vertebral column’s anatomic integrity is readily imaged with MRI and CT14, 15. Although retrospective clinical studies16–18 reported defect geometry, destruction of the pedicles, pain, age, anatomic site, lesion type, and activity levels to be risk factors of PVF, clinical guidelines for predicting PVF remain subjective with low specificity and sensitivity19. We postulate this poor performance in part due to the protocol’s failure to identify the specific features that result in decreased spinal stability20 and spinal strength6–8, 21–23. We6, 21, 24, and others7, 8, 23, 25, 26, have documented the effect of osteolytic lesion size, shape, location within the vertebral body7, 25, and cortical involvement7, 21, 25, on the vertebral strength. However, the biomechanical mechanisms underlying the association of mechanical instability with the pathological spine loss of strength remains unclear.

This in-vitro study simulated osteolytic defect models involving either the anterior or the anterior + middle vertebral columns in 3-level thoracic and lumbar cadaveric spines to investigate the association between the resulting kinematic instability and the loss of the lesioned spine’s compressive strength. We hypothesized that in spine segments with critical lytic defects, the loss of compressive strength is associated with the spine’s higher kinematical instability.

2. METHODS

2.1. Specimens:

As part of a previous study24, 3-level segments (T6-T8, T9-T11, T12-L2, and L2-L5) were obtained from eight fresh-frozen female cadaver spines, average (SD) age of 57(6.6) years [range: 47–69]. Each segment was CT-imaged (Aquilion 64, Canon Medical Systems, CA) to inspect for existing pathology or fractures and compute vertebral bone mineral density24. Twenty-four segments were selected for kinematic assessment (n=6/level), the segments dissected clean of musculature, and the cranial and caudal vertebrae shallow embedded in cement for mechanical testing24.

2.2. Preparation procedures: Figure 1 summarizes the experimental procedure.

Figure 1.

Figure 1.

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A pictorial illustration of the kinematic test protocol. First, we experimentally identified the segment’s instantaneous center of rotation (ICR) under applied compression, followed by an assessment of the segment’s kinematic in response to axial compression (AC), AC with flexion (AC+FL), and AC with extension (AC+EX) moments. This stage defines the segment’s baseline (control) kinematic response. Depending on the segment’s allocation, we mechanically created a large lytic defect model (Model C or model E) and repeated the kinematic assessment to evaluate the lesion’s effect on the segment’s kinematics. Lastly, we measured the lesioned segment compressive strength.

I). Preparation procedure.

To account for each segment’s unique kinematics and minimize coupled forces and moments imposed by test boundary conditions, we applied the test’s compressive loads at the segment’s center of force balance (CFB). Detailed in Appendix A.1, the segment’s CFB was identified (Figure 2), and its location about the segment’s mid-sagittal and mid-coronal axes marked on the cement embedding using indelible ink. These marks were used to co-register the segment between consecutive test stages. The segment’s middle vertebra was instrumented with a cluster of four passive optical markers (7mm, Qualysis AG, Sweden), and the segment secured to a six-degrees-of-freedom testing frame (Figure 3) with its mid-sagittal and mid-coronal axes aligned with the frame’s X and Y axes. As detailed in Appendix A.2, the testing frame was used to apply a 15N axial compressive (AC) follower load (i.e., a compressive load applied in line with the segment’s CFB), the segment allowed to equilibrate for 60sec under load, and the optical marker configuration acquired using cameras (ProReflex 1000, Qualisys AB, Sweden). This configuration establishes the marker- and anatomical-based vertebral coordinate systems for computing vertebral body motion about the segment’s anatomical axes27, 28. The measurement error for angles and translations were 0.2deg and 0.1mm, respectively.

Figure 2.

Figure 2.

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A diagrammatic illustration of the compression mechanical test device used for a) identifying the segment’s instantaneous center of rotation (ICR, section 2.2.I) and b) assessing its compressive strength, section 2.4. The hydraulic test system (DDC 4000, Interlaken, MN) applies axial compressive loading via a needle bearing onto the steel plate secured to the cranial vertebra. This arrangement allowed the spine to deform freely under the applied load. We used the mechanical X-Y table to adjust the compressive force point of application about the segment’s sagittal and coronal planes. By controlling the ratio of applied force to measured moment response, the ICR location is identified. We provide a detailed description of the method used to locate the ICR in Appendix A.1.

Figure 3.

Figure 3.

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A diagrammatic illustration of the six-degree-of-freedom (6DOF) mechanical test frame and the coordinate system convention used to present vertebral motion. The spine segment (not shown for clarity) is secured to the device via two aluminum rings (part #1) identical to the compression test device’s rings. An axial compressive loading system (parts # 2–7) utilizes calibrated static weights to apply a flower load-based compressive load (180N) to the spine. Independent of the application of axial compressive loading, as part of the moment loading system (parts #8–12), the second set of calibrated weights is used to apply pure flexion or extension moments with a magnitude of 7.5Nm. The applied force and moment and the force and moment response of the spine measured via 6DOF load cell (#13, MC5–5000, AMTI, MA). We provide a detailed description of the device in Appendix A.2. Optical markers, secured to aluminum rings (#16), were employed to compute the caudal- and cranial- most vertebrae motions.

2.3. Kinematic assessment:

AC was increased to 180N, the segment allowed to equilibrate for 60sec, and the cameras used to re-acquire (25Hz, 3 sec) the optical marker configuration. The applied compressive force and resultant forces and moments were measured via a load cell (MC5–5000, AMTI, MA) recorded at 10Hz using LabView (2015, National Instruments, TX). At the end of the test, each force and moment response curve was processed (Labview) to inspect whether a 5% instantaneous change had occurred during equilibrium. We previously found such a change to indicate a structural failure in vertebrae with lytic defects6. If no failure was detected, either combined AC + pure flexion (AC+FL) or AC + pure extension (AC+EX) moment (7.5Nm) was applied, the segment allowed to equilibrate (60sec), and the marker’s configuration re-acquired. If no structural failure was detected, the loading sequence repeated using the opposite moment (Figure 1).

2.4). Creation of lytic defect models:

As part of a previous study24, we simulated two models of solitary osteolytic defects, reported being associated with increased PVF risk1. The first model, created in 13 segments, simulated anterior column compromise encompassing 40% of the vertebral body (Taneichi1 Model C, Figure 4). The second model (Taneichi1 Model E), created in additional 11 segments, extending model C to include the ipsilateral pedicle and facet joint, simulating anterior + middle column compromise (Figure 4). Appendix (A.3) details the method of simulation. As part of this study, the lesioned segments’ kinematic response was re-measured, section 2.2.II.

Figure 4.

Figure 4.

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Illustration showing the location and extent of the two lytic models employed and corresponding computed tomography (CT) images of the models created. Type C involves approximately 40% of the vertebral body cross-section. Type E = Type C+ the destruction of the ipsilateral pedicle and facet joint.

2.5). Mechanical test to failure:

The testing was detailed as part of a previous study24. In brief, the test set-up (Figure 2) imposed a monotonically increasing compressive displacement (5mm/min) at the segment’s CFB. The test was terminated when a 10% reduction in maximum compressive force was measured (10kN range, 0.01N accuracy, Interlaken, MN), indicating fracture. The segment’s compressive strength was defined as the maximum compressive force value measured from the force-displacement curve.

2.6). Motion analysis:

The optical measurement files were processed28 to compute the control and lesioned segments range of motion (ROM), computed about the primary axis of rotation, and translations, computed along the main anatomical axes, about the superior vertebrae. Figure 3 details the kinematic definition used for the measurements.

2.7). Statistical analysis:

statical analysis was performed using JMP (11.0, SAS Institute, Cary, NC). Univariate analysis showed the motion parameters (translations, rotations) to be normally distributed (Shapiro Wilk test, p>0.05, respectively). For each test (AC, AC+FL, AC+EX), a two-way repeated measure analysis was used to test for the lesion model (C, E) effect on the change in motion parameter (translations, rotations) between the control vs. lesion spines. The sampling of multiple segments per spine can introduce non-independence of the data29. We fitted linear mixed-effects models under different assumptions about the correlation structure among segments from the same spines to test this effect. Based on the Akaike information criterion, the independence structure best fits the data29. We used Wilcoxon signed-rank to test the lesion model effect on spine strength.

Linear regression analysis was used to test whether the change in lesion mediated kinematic parameters, computed as [((lesion-intact) / intact)*100]%, was associated with the change in the segment’s compressive strength. Multivariable linear regression analysis was used to test the independent contribution of lesion model (C, E), age, and BMD to the kinematic parameter’s prediction of the segment’s compressive strength. All reported p values are two-tailed with an alpha level set to 0.05.

3. RESULTS

No specimen fractured during kinematic tests, either pre or post-introduction of the lesion models. For the following sections, C: and E: denote models C and E.

3.1. Lesion affect the segment’s kinematic response.

Descriptive statistics revealed that vertebral ROM and translation motions were non-normally distributed (Shaprio-Wilk test, P=0.0004 to 0.0456). Table A.1 details median (q1-q3) values for the control and lesion segments’ kinematic response. In comparison to the control segments:

  • AC loading: Model C affected higher sagittal translations (p=0.0162), Figure 5. Model E affected higher flexion (p<0.0001), lateral bending (p<0.0001) and torsional (p=0.0164) ROM, and higher sagittal (p=0.0008), transverse (p<0.0001) and axial (p<0.0001) translations, Figure 5. Two-way ANOVA revealed model E to affect significantly higher change in the lesioned segments flexion (p<0.001) and lateral bending (p=0.005) ROM and sagittal (p=0.0490), transverse (p<0.001), and axial (p=0.0413) translations.

  • AC+FL loading: model C caused higher flexion ROM (p<0.0001) and axial translations (p=0.0002), Figure 6. Model E caused higher flexion (p=0.0007) and torsional (p=0.0324) ROM, and higher sagittal (p=0.0068), transverse (p=0.006) and axial (p=0.0010) translations, Figure 6. Two-way ANOVA revealed model E to affect significantly higher change in the lesioned segment’s sagittal translations (p=0.0452).

  • AC+EX: model C caused lower extension ROM (p<0.0001) but higher sagittal (p=0.0480), transverse (p=0.0017) and axial (p=0.0020) translations, Figure 7. Model E caused lower extension (p<0.0001), higher lateral-bending (p=0.0123) and torsional (p=0.0123) ROM, and higher sagittal (p=0.041), transverse (p=0.0010), and axial (p=0.0020) tranlsations, Figure 7. Two-way ANOVA analysis revealed model E to affect significantly higher change in the lesioned segment’s extension (p=0.0233) and torsional (p=0.0429) ROM and sagittal (p<0.001) translations.

Figure 5.

Figure 5.

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In response to applied axial compression (180), the lesion segments demonstrated significantly higher translation motion along all three anatomical axes than the control segments. The additional disruption of the middle column caused a significant increase in the segment range of motion and an additional increase in the translation motion, indicating a higher degree of destabilization.

Figure 6.

Figure 6.

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Independent of the lesion model, the lesion segments exhibited significantly higher flexion ROM and axial translation than the control segments under the AC+FL test. The additional disruption of the middle column (Model E) caused a higher degree of destabilization, indicated by the increase in torsional ROM and the sagittal and transverse based translation compared to model C effect kinematic response of the control spines.

Figure 7.

Figure 7.

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Independent of the lesion model, the lesion segments exhibited a reduction in primary ROM (extension) but significantly higher translation in all three anatomical axes than the control segments in response to the AC+EX test. The additional disruption of the middle column caused a higher degree of lateral bending and torsional ROM and the higher magnitude of change in translation motions in three anatomical axes compared to the effect of model C suggesting a higher degree of destabilization.

3.2). Spine strength:

Descriptive statistics revealed compressive strength values to be non-normally distributed (Shapiro-Wilk, p=0.0394), yielding a median, (q1-q3) value of 1757 (1176 – 2576)N, Table 1. There was no significant difference in compressive strength between Model C and Model E (Wilcoxon 2-sample test, p=0.119). Multivariable linear regression analysis incorporating lesion model, age, and BMD, found the lesion model significantly associated with strength (R2=17, p=0.0484). Neither age nor BMD was significantly associated with strength.

Table 1.

Specimen details and results for the compressive strength test.

Level Lesion model Age (years) BMD (g/CM2) Compressive strength (N)
T6-T8 Model C 51.3 (3.8) 0.40 (0.07) 1608 (785)
T6-T8 Model E 64.0 (4.4) 0.42 (0.04) 2900 (1811)
T9-T11 Model C 51.3 (3.8) 0.40 (0.04) 1853 (1426)
T9-T11 Model E 62.7 (5.7) 0.41 (0.05) 2689 (1661)
T12-L2 Model C 57.0 (11.1) 0.38 (0.04) 1048 (257)
T12-L2 Model E 57.7 (4.5) 0.42 (0.06) 2685 (1343)
L3-L5 Model C 58.7 (4.9) 0.42 (0.05) 2389 (239)
L3-L5 Model E 53.3 (5.8) 0.41 (0.05) 2101 (1735)
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Values are presented as mean (standard deviation). BMD: Bone mineral density.

3.3). Loss of spine compressive strength is associated with kinematic changes:

  • AC loading: Neither the difference in ROM nor translation parameters between the lesion spine and control spines were significantly associated with the segment’s compressive strength, Figure 8.

  • AC+FL loading: Higher differnce in flexion ROM [C: (R2=0.38, p=0.0243); E: (R2=0.42, p=0.0285)], sagittal translations [C: (R2=0.48, p=0.0091); E: (R2=0.57, p=0.0071)], and for model E, higher axial (R2=0.69, p=0.0016) and transverse (R2=0.40, p=0.0206) translations, were associated with lower segment’s compressive strength, Figure 9.

  • AC+EX loading: Higher difference in extension [C: (R2=0.79, p<0.001); E: (R2=0.83, p<0.001)] and torsional (E: R2=0.32, p=0.0011) ROM and higher axial (C: R2=0.44, p=0.0134; E: R2=0.89, p<0.001), sagittal (E: R2=0.46, p=0.0216) and lateral translations [C: (R2=0.38, p=0.0429); E: (R2=0.4, p=0.0367)], were associated with lower segment’s compressive strength (Figure 10).

Figure 8.

Figure 8.

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Present the association between lesion models C and E and the change in the lesioned segments’ kinematic response under the applied axial compression (180N) and the segment’s compressive strength. Independent of the lesion model, segments demonstrating the least change in either ROM or translations showed a higher compressive strength.

Figure 9.

Figure 9.

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In response to combined axial compression (180N) and flexion moment (7.5Nm), the change in flexion ROM and axial and sagittal translation, caused by the disruption of the anterior column (model C), were negatively associated with the segment’s compressive strength. The additional disruption of the segment’s middle column worsens the degradation of the affected level compressive strength.

Figure 10.

Figure 10.

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In response to combined axial compression (180N) and exertion moment (7.5Nm), lower extension ROM and higher translation in all three anatomical axes demonstrated significant negative association with compressive strength in segments with disruption of the anterior column (model C). Similar to the previous tests, the additional disruption of the segment’s middle column (model E) yields higher degradation of the affected level compressive strength. The significant association between the higher torsional ROM and higher transverse translation with loss of compressive strength suggests the ipsilateral facet joint’s loss destabilizes the segment, further exposing the vertebral structure to additional shear due to the asymmetry of loading.

Multivariable analysis revealed the additional destruction of the middle column (p<0.0001), higher axial translation (AC+FL: p=0.0002; AC+EX p=0.0021), and for AC+FL higher sagittal translations (p=0.0009), and lateral bending ROM (p=0.0116), to be significant independent correlates of lower lesion segments compressive strength, Table 2. Neither age nor BMD was significant independent correlates of the lesion segments’ compressive strength.

Table 2.

Multivariable model for predicting the compressive strength of spine segments.

Variable B SE B CI: 95% P
Applied Compression (180N) with Flexion (7.5) Nm
Lesion model −939.2 138.7 −1229.5, −649.0 <.0001
Øy (% difference) 2.5 0.9 0.6, 4.3 0.0116
Δx (% difference) −5.7 1.4 −8.7, −2.6 0.0009
Δz (% difference) −8.7 1.9 −12.7, −4.7 0.0002
Applied Compression (180N) with Extension (7.5) Nm
B SE B CI: 95% P
Lesion model −868.2 98.7 −1074.2, −662.2 <.0001
Øx (% difference) 43.2 9.9 22.4, 64.0 0.0003
Δz (% difference) −6.4 1.8 −10.2, −2.6 0.0021
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% difference: The difference in the measured kinematic parameter = [((Defect – Intact)/ Intact)*100] %.

Kinematic parameters: Øx: lateral bending angle, Øy: Flexion/extension angle, Δx: Sagittal translation: Δz: Axial translation.

Statistical parameters: B: Model coefficient, SE: Standard error, CI: Confidence interval, P: Statistical significance.

4. DISCUSSION

This study demonstrated that kinematic instability caused by lytic disruption of the anterior vertebral column degrades the spine’s structural rigidity, causing the reduction of spinal compressive strength independent of age or BMD. Disruption of the middle column enhanced both effects, affirming the added risk for vertebral fracture associated with middle column involvement30. These novel findings highlight the importance of quantitatively assessing the lytic spine structural rigidity to better predict fracture risk.

In the biomechanics literature, ‘instability’ is defined as abnormal kinematics under load31. The effects of critical lytic defects on the kinematics of lumbar and thoracic spines have not been reported. Our findings demonstrate lytic disruption of the anterior column affects significantly higher flexion ROM under AC and AC+FL, lower extension ROM under AC+Ex, and higher axial and transverse translations under these loading modes. Compared to model C, disruption of the middle column (model E) caused a significantly higher degree of translational motion in all three anatomical planes and higher torsional and lateral bending ROM under AC+FL and AC+EX. These findings, indicating increased asymmetry in the segment’s motion response, affirm the importance of evaluating lytic destruction of the pedicles and facet joints in the lumbar and the thoracic spine when evaluating impending instability in these regions1, 30.

In patients with spinal metastases, mechanical instability directly affects the risk of PVF13. Derived from retrospective patient data26 and cadaver spines7, 23, 26, computational models of spinal segments with lytic anterior column disruption predicted higher axial deformation of the vertebral body to increase spinal fracture risk under compression7, 23, 26 and bending moments25. Our finding of that lower spine compressive strength was associated with higher axial translations due to lytic disruption of the anterior column (model C) in extension but not flexion moments, partially supporting these model’s predictions.

We found middle column disruption (model E) to result in a moderate to strong association between higher axial translations and lower spine compressive strength under AC+FL and AC+EX tests. These findings support the clinical emphasis for lytic destruction of the posterior elements as a strong prognostic indicator for impending risk PVF. As indicated by the multivariable analysis, higher sagittal translations, and, for model E, additional transverse translations, were significantly associated with lower spinal compressive strength. This novel finding suggests shear-based deformation within the vertebral body degrades the lesioned vertebrae structural rigidity, a finding supported by previous computational studies7, 21. The development of non-linear strains within the cancellous bone8, 21, and buckling of the vertebral cortex7, 8, 21, 25, underlying this structural degradation, may explain the higher degree flexion ROM caused by physical deformation of the vertebral body.

Similarly, the observed higher vertebral body deformation under extension moment may result in early engagement of the remaining facet causing increased torsional ROM and sheer based transverse translations. We posit that combined, these loads cause increased deformation of the vertebral body in the sagittal plane, reducing extension ROM. We found no significant association between BMD and the lesioned segments’ strength, a finding in agreement with previous in-vitro32, 33, and in-vivo34 studies. Although a gold standard for estimation of vertebral strength in osteoporosis patients35, BMD is region-specific36, “structurally simplistic” measure, less able to account for the effect of highly non-linear strains and stresses in the remaining cancellous bone7, 21, 25 on its ability to support the strength of lesion vertebrae6.

This study has several limitations beyond the limitations of cadaveric spine use. Our cadaveric spines, selected from females aged 47–69, demonstrate osteoporosis based on computed BMD values24. Accelerated bone loss is a common comorbidity in breast cancer patients37 and prostate patients undergoing androgen deprivation therapy38. Although lytic lesions’ effect on the bone’s material properties remains unclear, pre-clinical39, 40, and cadaveric41, 42 studies found little difference between osteolytic and healthy vertebral bone. Our segments did not possess existing lytic foci or permeative lytic lesions; instead, we created simulated lytic lesions to standardize the size and locations of defects. Thus, this study cannot fully reflect this patient cohort.

Spinal motion in vivo is complex. To account for segment-specific kinematics43, we elected to apply a follower-load compressive loading at the segment’s CFB. However, CFB location dynamically changes during segmental motion causing a “corrective” moment by the applied compressive load. Although < 5% of the applied moment, we cannot discount an effect on the segments’ kinematics. The exclusion of the rib cage increases the segment’s ROM and decreases stiffness44. We would expect this omission to yield higher motion values for the thoracic segments in this study despite the added load-carrying capacity by a follower load44. We would not expect any effect on spine strength. Whether these differences affect the association between kinematic instability and spine strength is unclear. Finally, our protocol application of “static” loading cannot fully capture the dynamic motion of the segment’s range of motion. However, the method employed is experimentally straightforward, addressing the core questions of kinematic change under load in the presence of defects. Since we evaluated the same spines before and after introducing the lesions, the kinematic response change is appropriately attributed to the defects.

Conclusions:

In patients with osteolytic spine lesions, movement-related bone pain45, 46, assumed to be caused by spine instability11, 12, is a risk factor for PVF. Our findings of greater kinematic abnormality with lower strength may explain these clinical observations13, 30. Although direct radiologic assessment of kinematics is not clinically feasible at present, elucidating the role of aberrant kinematics in PVF pathophysiology will help guide treatment decisions and optimize management strategies for his important patient population.

Supplementary Material

Apendix
Table A.1

CLINICAL RELEVANCE:

This experimental study demonstrates lytic foci to degrade the spinal kinematic stability, directly affecting spinal compressive strength. Understanding the mechanisms for this degradation will help guide treatment decisions that address inferred instability and fracture risk in patients with metastatic spine disease.

Source of Funding:

This study was supported by the National Institute of Arthritis and Musculoskeletal and Skin Diseases (AR055582) and the Harvard Catalyst (UL1 TR002541).

Footnotes

Investigation performed at Beth Israel Deaconess Medical Center (BIDMC) and Harvard Medical School, Boston, Massachusetts

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Supplementary Materials

Apendix
NIHMS1810659-supplement-Apendix.docx (24.3KB, docx)
Table A.1
NIHMS1810659-supplement-Table_A_1.docx (18.4KB, docx)

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