While both patient and cervical flexion—extension device were inside a clinical 3. The flexion—extension protocol was repeated five times in series over three separate sessions. To simulate the change of subjects, a 1 h break was taken between sessions during which the MRI bed was returned to the home position where the phantom was removed from the loading device and subsequently returned to its previous position using registration markers.
The resultant displacements were then averaged and evaluated for precision Supplementary Information. Displacement-encoded MRI was accomplished via a DENSE displacement encoding with stimulated echoes pulse sequence combined with a ssFSE single-shot fast spin echo sequence 32 , 70 synchronized with pneumatically-actuated cyclic flexion—extension of the neck.
DENSE displacement encoding was completed with a 0. DENSE phase map acquisition was performed in both anterior-posterior x -axis and cranio-caudal y -axis directions. Potential subjects were pre-screened for prior neck or back injuries through interview and those with signs of asymptomatic morphological abnormalities in the IVD during the preliminary MRI examinations were excluded. To minimize inter-subject variability and establish a consistent focal point of bending about the C7T1 region, fluid capsules served as fiducial markers—one placed onto the hinge of the loading platen and the other marker placed on the C7 spinous processes.
The subject was adjusted as necessary until the two markers were in a maximum proximity and within the localizer sequence FOV. To minimize body movement, subjects breathed and swallowed in sync with the non-acquisition period 4 s of the 8 s loading cycle. A single loading sequence consisted of a flexed state time of 2 s, a transition period of 0. After the preconditioning cycles, the dualMRI sequence was performed. To achieve a sufficient signal to noise ratio SNR i.
Total scanning time was approximately 45 min per subject. Therefore, analysis was focused on the cervical spine to specifically evaluate strains during simple flexion—extension i. For bulk analysis, the displacement, strain, and relaxometry values for the ROIs were averaged for each IVD and then averaged across all participants Figs.
Spatial analysis Fig. Voxel values were combined for all sections and then averaged. Data sets were evaluated for normality by a Shapiro—Wilks test. Sets which failed a Shapiro—Wilks test were transformed to meet the normality assumption by using either a square root transformation, in the case of non-negative data sets, or the following equation for data sets with negative numbers:.
The relationship between strain and relaxometry data Fig. Yeh, W. Elastic modulus measurements of human liver and correlation with pathology. Ultrasound Med. PubMed Article Google Scholar. Murphy, M. Imaging 34 , — Laklai, H. Genotype tunes pancreatic ductal adenocarcinoma tissue tension to induce matricellular fibrosis and tumor progression.
Tomasek, J. Myofibroblasts and mechano-regulation of connective tissue remodelling. Cell Biol. Karamichos, D. Collagen stiffness regulates cellular contraction and matrix remodeling gene expression. Part A 83A , — Adams, M. What is intervertebral disc degeneration, and what causes it?. Spine Phila Pa 31 , — Article Google Scholar. Nightingale, T. A model of unloaded human intervertebral disk based on NMR relaxation.
Driscoll, T. The global burden of occupationally related low back pain: estimates from the Global Burden of Disease study. Article PubMed Google Scholar. Murray, C. The state of US health, — JAMA , Todd, A. Cervical spine: degenerative conditions. Hogg-Johnson, S. The burden and determinants of neck pain in the general population. Spine Phila Pa 33 , S39—S51 Intervertebral disc degeneration: evidence for two distinct phenotypes.
Dudli, S. Pathobiology of modic changes. Spine J. Guterl, C. Challenges and strategies in the repair of ruptured annulus fibrosus. Cells Mater. Iatridis, J. Role of biomechanics in intervertebral disc degeneration and regenerative therapies: what needs repairing in the disc and what are promising biomaterials for its repair?. Setton, L. Mechanobiology of the intervertebral disc and relevance to disc degeneration.
Bone Jt. Google Scholar. Raj, P. Intervertebral disc: anatomy—physiology—pathophysiology-treatment. Pain Pract. Menezes, N. T2 and T1rho MRI in articular cartilage systems. Chan, D. Probing articular cartilage damage and disease by quantitative magnetic resonance imaging. Interface 10 , Rajasekaran, S. ISSLS prize winner: a study of diffusion in human lumbar discs: a serial magnetic resonance imaging study documenting the influence of the endplate on diffusion in normal and degenerate discs.
Spine Phila Pa 29 , — Kim, W. Application of elastography for the noninvasive assessment of biomechanics in engineered biomaterials and tissues. Streitberger, K. In vivo multifrequency magnetic resonance elastography of the human intervertebral disk. Walter, B.
MR elastography-derived stiffness: a biomarker for intervertebral disc degeneration. Radiology , — In vivo articular cartilage deformation: noninvasive quantification of intratissue strain during joint contact in the human knee. Neu, C.
Displacement encoding for the measurement of cartilage deformation. Articular cartilage deformation determined in an intact tibiofemoral joint by displacement-encoded imaging. Characterization of engineered tissue construct mechanical function by magnetic resonance imaging. Tissue Eng. Transient and microscale deformations and strains measured under exogenous loading by noninvasive magnetic resonance. Griebel, A. Noninvasive dualMRI-based strains vary by depth and region in human osteoarthritic articular cartilage.
Noninvasive assessment of osteoarthritis severity in human explants by multicontrast MRI. Mechanical deformation and glycosaminoglycan content changes in a rabbit annular puncture disc degeneration model. Spine Phila Pa 36 , — Intervertebral disc internal deformation measured by displacements under applied loading with MRI at 3T.
Comparison of intervertebral disc displacements measured under applied loading with MRI at 3. Johannessen, W. Assessment of human disc degeneration and proteoglycan content using T1rho-weighted magnetic resonance imaging.
Auerbach, J. Chen, C. Quantitative T2 magnetic resonance imaging compared to morphological grading of the early cervical intervertebral disc degeneration: an evaluation approach in asymptomatic young adults. Stelzeneder, D. Quantitative T2 evaluation at 3. Tropiano, P. Using a finite element model to evaluate human injuries application to the HUMOS model in whiplash situation.
Mwale, F. Quantitative MRI as a diagnostic tool of intervertebral disc matrix composition and integrity. Paul, C. Akamaru, T. Adjacent segment motion after a simulated lumbar fusion in different sagittal alignments: a biomechanical analysis.
Spine Phila Pa 28 , — Hilibrand, A. Adjacent segment degeneration and adjacent segment disease: the consequences of spinal fusion?. Siskey, R. Development of a clinically relevant impingement test method for a mobile bearing lumbar total disc replacement. Panjabi, M. Injury mechanisms of the cervical intervertebral disc during simulated whiplash. Teraguchi, M. Allied to these remarkable developments, numerical studies through finite element analysis FEA are benefiting from the improvements on the computational power to become more exhaustive and wide-ranging every day Schmidt et al.
The bioreactor system was developed within this group, and is denominated as loaded disk culture system LDCS. In short, this system is capable of maintaining an IVD alive for at least 3 weeks once extracted, after the sacrifice of the animal, which is usually a goat. In comparison with the abovementioned analogous bioreactors, this is the one, which reports the longest period of IVD viability.
The loading feature allows the IVD to be submitted to compression tests without losing its biomechanical and physiological properties Paul et al. A schematic representation of this system is shown in Figure 1. Figure 1. Schematic representation of the LDCS. Adapted from Paul et al. The LDCS reproduces the IVD native environment, by enabling the close monitoring of oxygen and nutrient supply levels, through the introduction of a culture medium, along with providing static or dynamic mechanical axial loading, as seen in Figure 1.
The previous LDCS-related publications reported that, from the physiological point of view, a dynamic loading regime cause large dynamic displacement, while the static regimes induced a prolonged creep effect Paul et al.
Up to 12 IVDs may be simultaneously under experiment in this equipment. The object of study is an LDCS dataset, which contained the results of experiments performed with eight partial lumbar MS, i. The CABC is an enzyme that cleaves proteoglycans, i. The lumbar IVDs used in the current study are derived from goats, which were also used in the study published by Detiger et al. Although the individual IVDs were not used for analyses in that study, the entire lumbar spinal segments of the goats were scanned with MRI pre-operatively and 12 months after CABC injection.
Detiger and co-workers reported a large inter-individual variation for all parameters measured. However, no significant difference in Pfirmann scale was measured for the IVDs included in the current study data not previously published. The partial MS were tested under physiological loading conditions, which consisted of a sinusoidal load 1 Hz of N average and N amplitude for 16 h activity period , followed by other sinusoidal load 1 Hz of 50 N average and 10 N amplitude for 8 h resting period.
Sinusoidal loadings are associated with dynamic loading regimes Qasim et al. This loading profile is comparative to activities such as lying down and walking in goats and relaxed standing and non-supported sitting in humans. It must be highlighted that the transition between the activity and resting periods is performed with 1 h of triangular loading 0. The most relevant material properties of this model are summarized in Table 1 , accordingly to the state-of-the-art of soft tissue and IVD constitutive modeling Dreischarf et al.
Table 1. Material properties of the partial physiological MS FE model. Figure 2. Sagittal cut of the partial human L3—L4 MS FE model, which contains node quadratic hexahedral elements and 16, nodes. The average axial cross section of the goat IVD is around — mm 2 Paul et al. In what concerns to the IVD height, the FE model has an average height of 12 mm, while the goat IVDs have an average height of 9 mm, which means that no normalization is needed.
Previous studies verified that human and goat IVDs produce similar internal stresses, regardless of the geometric differences Ayotte et al. The other major simplification is related with the loading regimes. At the hourly time-scale, it was numerically verified that the sinusoidal loadings applied in the LDCS and the equivalent static average loadings produce comparable numerical outcome loadings.
Since the sinusoidal wave meant additional computational effort, one adopted the lighter static and constant loadings. Two daily cycles were simulated, after a preconditioning period of 8 h, for osmotic equilibration before loading. The numerical physiological loading profile consists of N for 16 h activity period , followed by N for 8 h resting period. The transition between the activity and resting periods is performed with N for 1 h, in order to maintain a similar timeline with the experimental tests.
As aforementioned, CABC cleaves proteoglycans, which are involved in the osmotic pressure mechanisms, through the regulation of the fixed charged density C F. Hence, if the proteoglycan content is reduced, the hydration of the IVD is also reduced and the osmotic pressure gradient is lower Wognum et al. The initial fixed charge density C F ,0 of the NP was then reduced, in two levels, until it reached the initial fixed charge density of the AF, which remain unaltered Table 1.
This dual-step reduction is presented on Table 2. Table 2. Osmotic swelling material properties and correspondent initial osmotic pressure of both native and reduced OsmP FE models. Mild degeneration, as reported by Detiger et al. Low to mild degeneration are probably related to a decrease of CEP permeability, due to the calcification of this component, while mild to severe degeneration are more likely associated with an increase of CEP permeability, as a result of probable crack openings Urban, ; Adams and Dolan, ; Stefanakis et al.
The permeability variations on the degenerated NP and AF remain uncertain, but some studies pointed out that permeability could increase Iatridis et al.
The proteoglycans might also be related to the permeability variations, but such association is also unclear. The IVD components are also described to increase in stiffness, as modeled by Natarajan et al. The timeline was limited to two daily cycles. Other IVD-related outputs could be extracted from the FE solver, namely intradiscal pressure, fiber elongation of volume variation. Figure 7 shows the comparison between the injected IVDs DHV results of the experimental and the corresponding numerical tests, i.
Table 3 summarizes the DHV values of these seven DHV curves, demonstrating the differences between the four experimental tests, and also between the experimental tests and the numerical ones. Figure 3. Figure 4. Figure 5. Figure 6. Native DHV outcomes of the experimental and numerical tests, for two daily cycles, i. Figure 7. Non-physiological DHV outcomes of the experimental and numerical tests, for two daily cycles, i. Table 3. Summary of the DHV values of the four injected IVDs and the reduced OsmP models, after each activity and recovery period, obtained by numerical simulation.
The control IVDs tend to maintain their height along the experiment, while the injected IVDs tend to lose height progressively Figures 3 and 4 , respectively. This systematic height reduction is mostly noticed on the loss of the ability to return to the same height level after the recovery period.
Such findings are in agreement with the work of Detiger et al. However, some discrepancies were noticed between the injected IVDs, which can be justified by the inter-specimen variability of the goats. In fact, goats 1 and 3 seem to be more sensitive to the CABC compound than goats 2 and 4.
In what concerns to Figure 5 , it is noticeable that the behavior of the numerical model corresponds to what was theoretically expected. It must be highlighted that the transition periods are excluded from the current analysis of the maximum and minimum displacement measurements, as their purpose was just to establish the transition between the activity and recovery periods, on the experimental apparatus. Secondly, the reduced OsmP models show some signs of degeneration, as the fluid flow seems to be diminished, and thus the initial height is not recovered.
The recovery rate shall be understood as the difference between the DHV after the activity period and the DHV after the recovery period.
Considering the described outcomes, the situation of an IVD without any osmotic pressure gradient is certainly related to severe degeneration Adams et al. In what concerns to the resting periods, an important difference is noticed, as the numerical model is able to regain all the fluid lost during the activity period. Nonetheless, these four control IVDs maintain the same DHV recovery level from the first to the second daily cycle, which was previously described as a sign of no degeneration.
In other words, the DHV results indicate no degeneration, but incomplete recovery. This fact is probably related with the intrinsic behavioral differences between the goat and human IVDs, namely the specific biomechanical stimuli. The nutrition pathways included in the LDCS functioning and the ion-influenced fluid retention may also have a role in this situation. The numerical model helped to understand that the ideal situation is to fully recover the fluid on the resting periods.
However, the action of AF fibers may also be limiting the range of DHV, and this limit situation maximum extension of the fibers is not predicted in the FE solver. This action shall not be considerably different in human and goat spines. In addition, the analogous difference for the maximum displacement is averagely 0. The DHV measurements from the injected IVDs of goats 2 and 4 may not be directly compared with these numerical outcomes, as their maximum displacement is significantly lower than the analogous measurement of the other two injected IVDs.
The DHV measured after each recovery period is nearer to the numerical models, but the recovery rate is considerably uneven. On the numerical point of view, the results suggest that the MS FE model is excessively sensitive to the applied loads, having that the goat IVDs present a much more limited range of height variation.
If these experimental tests were performed in vivo , one might question the amount of influence of the surrounding tissue ligaments on IVD mechanical properties and measurements. On the one hand, one shall not neglect the influence of the external membranes covering each MS, since they also play a role on limiting the range of motion of IVD and other organs Humphrey, On the other hand, previous studies from Ayotte et al. Eight spring elements were applied around the L1—L5 lateral vertebral body where the total load was divided equally to each of the spring element [ 24 , 25 ].
This is to assure the uniformity of the applied follower load and to avoid any potential rotation of the intervertebral body. The inferior surface of the L5 vertebral body was completely fixed in all directions as shown in Figure 2. The FE model of osseoligamentous lumbar spine was verified by comparing the range of motion ROM with previous in vitro study for flexion and extension motions at pure moment of 7.
The present results of the intersegmental rotations of the lumbar spine follow similar trend to the previous in vitro results as shown in Figure 3 [ 21 ]. The percentage difference of the ROM between present and previous in vitro study in flexion was 7. It was found that similar trends were observed in the previous in vitro studies as shown in Figure 4 [ 26 , 27 ]. Based on these results, the developed FE model could produce appropriate and reliable results for further FE analysis.
It was found that the IDP was increased as the human spine physiological loading increased in flexion motion where the highest pressure was 1. The IDP was increased in flexion motion but an opposite trend was observed in extension motion. The effects of the human weight were observed to be more significant at the L4-L5 segment as shown in Figure 6.
In general, the annulus stress increased as the human weight increased with the maximum annulus stress of 3. The present study demonstrates that physiological loading of body weight plays an important role of stress distribution at IVD in the lumbar spine. It was observed that increasing body weight will increase the pressure at nucleus pulposus and annulus fibrosis at all levels of the IVD. Furthermore, the position and direction of motion appears to affect these results where the IDP was increased in flexion motion but an opposite trend was observed in extension motion.
A severe effect was noticed when heavier individuals continue to experience increased stress and pressure on IVD at all vertebral segments in the lumbar spine particularly at L4-L5 segment in both flexion and extension motions. The results of the present study show similar pattern to the IDP measured in in vitro study where the maximum IDP was found in flexion motion due to the load shift from the posterior towards the anterior of the IVD in flexion motion [ 12 , 28 , 29 ]. The increase of nucleus pressure enhances the tensile stress on the annulus fibers which leads to the excessive stress on the IVD and stimulate the propagation of disc degeneration particularly in nucleus pulposus [ 21 ].
This will increase the stiffness of IVD and could reduce its height due to the outflow of fluid through the vertebral body endplates [ 30 ]. Subsequently, the fluid loss will increase the proteoglycan and osmotic pressure within the nucleus [ 30 , 31 ]. The maximum annulus stress was obtained in extension motion due to the load shift from the anterior towards the posterior of the IVD. This phenomenon was also observed in previous clinical study where the structural defect in the vertebral body endplate tends to distribute the load transferred from the nucleus to the posterior annulus.
It has been shown that this can potentially lead to pain and could tear the annulus at the disc rim [ 32 , 33 ]. Gaining body weight will increase stresses of IVD at all vertebral segments in the lumbar spine particularly the L4-L5 segment. Furthermore, the nucleus pulposus was more severely affected compared with the annulus fibrosus.
Although flexion and extension motions of the lumbar spine appear to have different percentage effect on the IVD, it was found that heavier individuals will continue to experience an increase in stress at IVD regardless of the position of the spine.
This could be a factor that can lead to early intervertebral disc damage. This is an open access article distributed under the Creative Commons Attribution License , which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. Article of the Year Award: Outstanding research contributions of , as selected by our Chief Editors.
Read the winning articles. Journal overview. Special Issues. Academic Editor: Lizhen Wang. Received 05 May Revised 10 Jul Accepted 24 Jul Published 27 Aug Abstract The present study was conducted to examine the effects of body weight on intradiscal pressure IDP and annulus stress of intervertebral discs at lumbar spine. Introduction Obesity has been recognised as a factor that could lead to chronic low back pain LBP.
Materials and Methods 2. Finite Element Modeling The geometrical data of lumbar vertebrae were obtained from computed tomography CT scan of a healthy year-old male with 1. Figure 1. L1—L5 lumbar spine. Table 1. Geometrical parameters of the lumbar spine ligaments [ 15 , 18 ].
Table 2. Material properties of the components in the osseoligamentous lumbar spine model. Figure 2. Loading and boundary condition of FE model of the lumbar spine. Table 3. Figure 3. Figure 4. Figure 5. Figure 6. IDP contour plots of nucleus pulposus at L4-L5 vertebral segment.
Figure 7. References P. Gordon-Larsen, H. Wang, and B. Djurasovic, K. Bratcher, S. Glassman, J. Dimar, and L. Vismara, F.
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