Biomechanical Changes of Adjacent Segments Following Single- versus Two-level Oblique Lateral Interbody Fusion with Different Fixation Methods: A Finite Element study

Biomechanical Changes of Adjacent Segments Following Single- versus Two-level Oblique Lateral Interbody Fusion with Different Fixation Methods: A Finite Element study | Research Square window.SnipcartSettings = { analytics: { enabled: false } }; (function() { var accessVector = localStorage.getItem('access_vector') || ''; window.dataLayer = window.dataLayer || []; if (accessVector) { window.dataLayer.push({ user: { profile: { profileInfo: { snid: accessVector } } } }); } })(); (function(w,d,s,l,i){w[l]=w[l]||[];w[l].push({'gtm.start':new Date().getTime(),event:'gtm.js'});var f=d.getElementsByTagName(s)[0],j=d.createElement(s),dl=l!='dataLayer'?'&l='+l:'';j.async=true;j.src='https://www.googletagmanager.com/gtm.js?id='+i+dl;f.parentNode.insertBefore(j,f);})(window,document,'script','dataLayer','GTM-K279D39R'); Browse Preprints In Review Journals COVID-19 Preprints AJE Video Bytes Research Tools Research Promotion AJE Professional Editing AJE Rubriq About Preprint Platform In Review Editorial Policies Our Team Advisory Board Help Center Sign In Submit a Preprint Cite Share Download PDF Research Article Biomechanical Changes of Adjacent Segments Following Single- versus Two-level Oblique Lateral Interbody Fusion with Different Fixation Methods: A Finite Element study Shuo Li, Hengrui Chang, Kaibin Fan, Lifei Wang, Di Zhang, Jianhua Ren, and 2 more This is a preprint; it has not been peer reviewed by a journal. https://doi.org/ 10.21203/rs.3.rs-8272176/v1 This work is licensed under a CC BY 4.0 License Status: Under Review Version 1 posted 14 You are reading this latest preprint version Abstract Background This study investigated the biomechanical effects of single- and two-level oblique lateral interbody fusion (OLIF) with different internal fixation methods on adjacent segments in an osteoporotic lumbar spine model, focusing on range of motion (ROM), intervertebral disc stress, endplate stress, and facet joint contact forces. Methods A three-dimensional finite element (FE) model of the L1–S1 spine was developed from CT scans of a healthy 30-year-old male and modified to simulate osteoporosis. The model included vertebral bodies, posterior elements, endplates, discs, and major ligaments. Material properties were assigned based on established literature, with modifications to simulate osteoporosis. Nine surgical scenarios were simulated: standalone cages, bilateral pedicle screws (BPS), and cortical bone trajectory (CBT) screws for both single- and two-level OLIF. A 400 N follower load and 10 Nm pure moments in six directions were applied to L1 to replicate physiological loading. Results All fusion constructs increased ROM and mechanical stress at adjacent segments compared with the intact osteoporotic spine. Two-level fusion induced significantly greater biomechanical alterations than single-level fusion. In the L2–L3 segment, the L3–L5 + BPS model exhibited the highest flexion range of motion (ROM) of 6.70° and peak disc stress of 2.20 MPa, whereas the L3–L4 cage + BPS configuration demonstrated a ROM of 6.07° and peak disc stress of 2.02 MPa. Similarly, at the L5–S1 level, the L3–L5 + BPS construct produced the greatest flexion ROM (14.69°) and disc stress (3.67 MPa), compared with the L4–L5 + BPS construct, which yielded a ROM of 13.14° and disc stress of 3.34 MPa. CBT fixation consistently produced lower disc and endplate stresses compared with BPS fixation; however, it resulted in substantially higher facet joint stresses, particularly during axial rotation. In the L3–L5 + CBT construct, L5–S1 facet joint stress increased by 143% relative to the intact condition. Conclusion Two-level constructs impose greater cumulative loads than single-level ones. Compared with bilateral pedicle screws, cortical bone trajectory screws reduce anterior column loading but increase facet joint stress, especially during rotation. lumbar fusion osteoporosis finite element analysis adjacent segment degeneration cortical bone trajectory screws pedicle screws biomechanics Figures Figure 1 Figure 2 Figure 3 Figure 4 Figure 5 Figure 6 Figure 7 Figure 8 Figure 9 Background Lumbar spinal fusion is widely employed for the treatment of degenerative and unstable spinal disorders. However, complications including implant failure, cage subsidence, and adjacent segment degeneration (ASD) remain prevalent, particularly in elderly patients with compromised bone quality. Osteoporosis, which is common in this group, weakens bone and alters spinal biomechanics, raising the risk of screw loosening and construct instability [ 1 , 2 ]. ASD is a major reason for reoperation and long-term disability after lumbar fusion, with 10-year incidence rates reported between 15% and 18% [ 3 – 6 ]. OLIF has gained widespread adoption in the management of degenerative lumbar disorders due to its capacity to restore disc height, correct sagittal alignment, and achieve high fusion rates with minimal soft tissue disruption[ 7 – 9 ]. While stand alone OLIF may provide adequate stability in select single level pathologies, multilevel constructs often require supplemental posterior fixation to mitigate the risk of cage subsidence and pseudarthrosis, especially in osteoporotic patients[ 10 , 11 ]. Compared with the standard fixation method of bilateral pedicle screws (BPS) for lumbar fusion, cortical bone trajectory (CBT) screws have been reported in recent years to provide similar pull-out strength and resistance to buckling in osteoporotic bone, with less soft tissue injury and faster recovery[ 12 – 16 ]. Moreover, some clinical data have shown that cortical bone trajectory (CBT) screws and bilateral pedicle screws (BPS) have similar fusion rates and patient-reported outcomes in single-level surgery[ 17 – 20 ]. Currently, there are various studies on the impact of single-level OLIF surgery on adjacent segment vertebrae, including clinical studies, in vitro biomechanical studies, and finite element studies[ 21 – 24 ]. To date, no study has systematically investigated how these techniques influence load distribution at both the superior and inferior adjacent segments across different fusion configurations, including single-level and two-level fusions, under physiologically relevant loading conditions.To address this, we developed and validated a detailed finite element model of the osteoporotic L1-S1 lumbar spine. This study aimed to investigate the biomechanical effects of single- and two-level OLIF with different internal fixation methods on adjacent segments in an osteoporotic lumbar spine model. Methods Construction of an Intact Lumbar Finite Element Model The finite element (FE) modeling approach used in this study is identical to that established by Fan et al[ 25 ]. from our research group for biomechanical analysis of osteoporotic lumbar spines following OLIF. Their study investigated the biomechanics of two-level OLIF in osteoporotic spines with different internal fixation techniques and found that supplemental fixation enhanced segmental stability and reduced cage stress. BPS outperformed unilateral pedicle screws (UPS) and CBT screws in restricting segmental motion and lowering stresses on cages and implants. Here, we applied the same FE model to evaluate the biomechanical effects of single- and two-level fixation combined with OLIF on adjacent segments in an osteoporotic spine. Briefly, a three-dimensional FE model of the intact L1-S1 lumbar spine was reconstructed from computed tomography (CT) scans of a 30-year-old healthy male volunteer (slice thickness: 0.625 mm; 570 slices in DICOM format). Segmentation and initial 3D geometry reconstruction were performed using Mimics 21.0 (Materialise, Leuven, Belgium), followed by surface refinement in Geomagic Studio 12.0 (3D Systems, Rock Hill, SC, USA). Intervertebral discs and other soft tissues were modeled in Creo 8.0 (PTC, Boston, MA, USA), and the final anatomical structures were meshed using ANSA (BETA CAE Systems, Thessaloniki, Greece). The model includes vertebral bodies (L1-S1), posterior bony elements (pedicles, laminae, facets), cartilaginous endplates, intervertebral discs (annulus fibrosus and nucleus pulposus), and seven major spinal ligaments (anterior longitudinal ligaments, posterior longitudinal ligaments, ligamentum flavum, supraspinous ligaments, interspinous ligaments, capsular ligaments, and intertransverse ligaments). A mixed mesh of tetrahedral and hexahedral elements was used, with an element size of 1 mm determined through convergence analysis (ROM variation < 1% between 1 mm and 2 mm meshes under 400 N compression + 7.5 Nm flexion). Material properties were assigned based on established literature [ 25 ]: cortical bone (E = 12.0 GPa), cancellous bone (E = 100 MPa), endplates (E = 4 GPa), and posterior elements (E = 3.5 GPa). To simulate osteoporosis, the elastic modulus of cortical, the endplates, and the posterior elements were reduced by 33%, and the cancellous bone were reduced 66% respectively [ 25 , 26 ]. The annulus fibrosus was modeled as a fiber-reinforced composite with 12 layers of collagen fibers oriented at ± 30°(E = 4.2 MPa), while the nucleus pulposus was treated as nearly incompressible (E = 0.1 MPa, ν = 0.49). Ligaments were represented as nonlinear tension-only truss elements, and facet joints were defined as surface-to-surface contact pairs with a friction coefficient of 0.1. The detailed material properties of each component are listed in Table 1 . Table 1 Material properties of the FEM and implants Components Young’s modulus (MPa) Poisson ratio Cortical bone 8040 (“normal”: 12000) 0.3 Cancellous bone 34 (“normal”: 100) 0.2 End-plate 2640 (“normal”: 4000) 0.4 Posterior elements 2345 (“normal”: 3500) 0.25 Nucleus pulposus 1 0.49 Annulus 4.2 0.45 Anterior longitudinal ligament 20 0.3 Posterior longitudinal ligament 20 0.3 Ligamentum flavum 19.5 0.3 Interspinous ligament 11.6 0.3 Supraspinous ligament 15 0.3 Intertransverse ligament 58.7 0.3 Capsular ligament 32.9 0.3 Cage 3500 0.3 Screw and rod 110,000 0.3 Boundary and Loading Conditions Boundary and loading conditions followed standard protocols: the inferior endplate of S1 was fully fixed (all six degrees of freedom constrained), a 400 N follower load was applied to the superior endplate of L1, and a pure moment of 10 Nm was subsequently applied in six directions (flexion, extension, left bending, right bending, and left rotation, and right rotation) to simulate physiological spinal motion. Surgical Models and Fixation Configurations Three-dimensional geometric models of the internal fixation devices were developed in Creo 8.0 using the Part module, based on the actual dimensions of the interbody cage and supplemental fixation components. The interbody cage was modeled after the Oracle cage (DePuy Synthes), measuring 40 mm in length, 22 mm in width, 11 mm in anterior height, 8 mm in posterior height, and featuring an 8° lordotic angle. The pedicle screws had a diameter of 6.5 mm and a length of 50 mm, while the cortical bone screws were 5.0 mm in diameter and 30 mm long. The connecting rods had a diameter of 5.5 mm. Both the cage and the supplemental fixation devices were discretized using tetrahedral meshing. Bone-screw interfaces were assumed to be fully bonded to simulate immediate postoperative stability. Detailed material properties for each component are provided in Table 1 . The surgical segment was defined as the L3–L5 intervertebral space, and the annulus fibrosus, nucleus pulposus, and cartilaginous endplate were removed from the left side. Nine surgical models were created to evaluate the biomechanical performance of different fusion strategies under osteoporotic conditions in Fig. 1 – 3 : Single-level L3-L4 fusion: L3-4 Cage: Standalone polyetheretherketone (PEEK) interbody cage; L3-4 BPS: Cage + bilateral pedicle screw system (6.5 mm diameter, 50 mm length); L3-4 CBT: Cage + bilateral cortical bone trajectory (CBT) screws (5.0 mm diameter, 30 mm length, 10° medial, 25° cephalad). Single-level L4-L5 fusion: L4-5 Cage, L4-5 + BPS, L4-5 CBT (same screw parameters) Two-level L3-L5 fusion: L3-5 Cage: Cages at both levels without posterior fixation; L3-5 BPS: Cages + bilateral pedicle screws at L3 and L5; L3-5 CBT: Cages + bilateral CBT screws at L3 and L5. Importantly, while the FE model and surgical configurations (including fusion levels, cage types, and fixation strategies) are identical to those in Fan et al. [ 20 ], the focus of the present study differs substantially. Whereas Fan et al. primarily investigated the biomechanical behavior of the instrumented (surgical) segments and internal fixation devices, this study specifically evaluates the biomechanical response of the adjacent non-fused segments, the cranial adjacent segment (L2-L3) and caudal adjacent segment ( L5-S1).Our research emphasis on range of motion (ROM) of adjacent vertebral segments, intradiscal stress distribution, and facet joint contact forces. These parameters are critical indicators for assessing the potential risk of adjacent segment degeneration (ASD) following lumbar fusion under osteoporotic conditions. Results Range of Motion (ROM) At the L2-L3 level, all instrumented models, demonstrated increased ROM relative to IntactOP model (flexion: 5.02°), confirming compensatory hypermobility due to increased caudal stiffness. The largest increases were observed in flexion (up to 6.70° in L3-5cage + BPS; +33.5%) and lateral bending (right bend: 6.26° vs. 5.23°; +19.7%). Similarly, at L5-S1, ROM consistently exceeded baseline values (IntactOP model flexion: 11.39°), with the greatest elevation seen in the L3-5cage + BPS model (flexion: 14.69°; +29.0%). These findings indicate that both proximal and distal adjacent segments experience significant kinematic overload following lumbar fusion. Two-level fusion (L3-5) consistently induced greater ROM increases at adjacent segments than single-level fusion (L3-4 or L4-5). At L2-L3, flexion ROM rose from 5.57° (L3-4cage) to 6.02° (L3-5cage; +8.1%), and further to 6.70°with BPS fixation (L3-5cage + BPS; +12.2% vs. L3-4cage + BPS). At L5-S1, L3-5 fusion models exhibited markedly higher motion than L4-5 counterparts: for example, flexion ROM was 14.69° (L3-5cage + BPS) versus 13.14° (L4-5cage + BPS; +11.8%). This trend held across all loading modes. Two-level fusion (L3-L5) resulted in even greater compensatory motion at both adjacent levels: L2-L3 ROM increased by 19.98% to 33.33%, while L5-S1 ROM increased by 15.20% to 28.99%, confirming that extending the fusion span amplifies adjacent segmental motion. Notably, the greatest ROM increases were observed in the BPS groups across all fusion types, whereas CBT fixation consistently demonstrated lower adjacent segment mobility compared to BPS. Figure 4 . Intervertebral Disc Stress Peak stresses in the intervertebral discs were elevated in all fusion models. Single-level fusion consistently elevated IVD stress compared to the intact condition, with posterior instrumentation exacerbating this effect. At the proximal adjacent segment (L2–L3), the L3–L4 cage-only construct increased disc stress to 1.79 MPa in flexion (a 16.2% increase from intact 1.54 MPa). The addition of bilateral pedicle screws (BPS) further raised stress to 2.02 MPa (+ 31.2%), while cortical bone trajectory (CBT) screws yielded an intermediate value of 1.95 MPa (+ 26.6%). A similar hierarchy was observed at the distal adjacent segment (L5–S1): L4–L5 cage increased stress to 3.06 MPa in flexion (intact: 2.75 MPa; +11.3%), whereas L4–L5 + BPS and + CBT reached 3.34 MPa (+ 21.5%) and 3.24 MPa (+ 18.2%), respectively. These trends were consistent across all loading modes, indicating that BPS fixation imposes the greatest additional load on adjacent discs, followed by CBT and cage-alone constructs. Extending fusion to two levels (L3–L5) further amplified adjacent disc stress beyond single-level scenarios. For example, at L2–L3, the L3–L5 + BPS model produced a peak stress of 2.20 MPa in flexion—18% higher than the 1.88 MPa observed with L3–L5 cage and 9% higher than L3–L4 + BPS (2.02 MPa). Similarly, at L5–S1, L3–L5 + BPS generated 3.67 MPa in flexion, exceeding both L3–L5 cage (3.32 MPa) and L4–L5 + BPS (3.34 MPa). Across all constructs, two-level fusion increased adjacent IVD stress by approximately 10–13 percentage points compared to anatomically matched single-level fusions, underscoring the cumulative biomechanical burden of longer constructs. Notably, the proximal adjacent segment (L2–L3) exhibited greater relative stress elevation than the distal segment (L5–S1) under identical fusion conditions. In the L3–L5 + BPS model, L2–L3 disc stress increased by 42.9% in flexion (from 1.54 to 2.20 MPa), whereas L5–S1 stress rose by 33.5% (from 2.75 to 3.67 MPa)—a difference of 9.4 percentage points. This pattern persisted across all two-level constructs and loading directions, with L2–L3 consistently demonstrating 7–12% higher relative stress increases than L5–S1. Despite the higher baseline stress in the intact L5–S1 disc (e.g., 2.75 MPa vs. 1.54 MPa in flexion), the proportional mechanical perturbation was more pronounced at the cranial adjacent level. (Fig. 5 ). The peak stress distribution in the L2–L3 and L5–S1 intervertebral discs under single-level and two-level surgical configurations as shown in Fig. 8 . Endplate Stress Endplate stress followed a pattern closely aligned with disc stress. At L2-L3, peak endplate stress increased by 20–52% in flexion and 19–45% in extension, reaching 13.82 MPa (L3-L5 + BPS) in flexion and 10.34 MPa in extension. At L5-S1, stress rose by 14–43% in flexion and 13–35% in extension, with the L3-L5 + BPS model again showing the highest values (16.28 MPa in flexion; 11.28 MPa in extension). The L3-L5 + CBT construct also induced substantial endplate loading (15.76 MPa in flexion), approaching that of BPS. Stand-alone cage fusion consistently demonstrated the smallest stress increases (13–20%) across all critical loading modes, significantly lower than constructs augmented with BPS or CBT screws (25–52% increase). Although CBT fixation yielded slightly lower stresses than BPS, it did not confer a substantial biomechanical advantage in mitigating adjacent segment loading. (Fig. 6 ) The peak stress distribution in the L2-L3 and L5-S1 endplate under single-level and two-level surgical configurations as shown in Fig. 9 . Facet Joint Stress In stark contrast to the anterior column, facet joint stresses were predominantly elevated during extension and rotation. At L2-L3, extension stress surged from 16.85 MPa (intact) to 26.02 MPa in the L3-L5 + CBT model (+ 54%), the highest value observed across all conditions. During rotation, L3-L5 + CBT also produced the greatest stresses (22.25 MPa left; 22.38 MPa right). At L5-S1, extension stress increased by 24–76%, with L3-L5 + BPS reaching 15.56 MPa (+ 76%). Most strikingly, during left axial rotation, the L3-L5 CBT construct generated an exceptionally high facet stress of 17.64 MPa, representing a 143% increase compared with the intact condition (7.25 MPa) and markedly exceeding the values observed in all other models. In contrast, facet stresses remained minimal during flexion (< 0.25 MPa at L2-L3; <4.08 MPa at L5-S1), confirming their non-load-bearing role in this motion. (Fig. 7 ). Discussion Adjacent segment degeneration (ASD) remains a leading cause of reoperation after lumbar fusion, with 10-year rates of 15% to 18% [ 3 , 5 ]. Biomechanical overload at adjacent levels—particularly the upper segment, where degeneration is consistently more severe [ 6 , 27 ]—is a primary driver, and this risk is amplified in osteoporotic patients due to poor bone quality and reduced screw–bone stability [ 1 , 2 ]. While cortical bone trajectory (CBT) screws offer improved pullout strength in low-density bone compared to conventional pedicle screws [ 12 , 14 ], their impact on adjacent segment loading in multilevel constructs remains poorly understood. Most finite element studies have focused on single-level fusion in non-osteoporotic models [ 28 – 30 ], limiting clinical relevance for elderly patients who often present with contiguous multilevel disease. To address this gap, we developed a validated osteoporotic L1–S1 finite element model to compare three fixation strategies—standalone interbody cage, bilateral pedicle screws (BPS), and CBT screws—under both single-level (L4–L5) and two-level (L3–L5) fusion scenarios, with L3–L5 selected as a representative multilevel degeneration pattern in older adults [ 25 ]. Model kinematics were validated against in vitro motion data [ 25 , 31 ] to ensure physiological fidelity. Our kinematic analyses confirm that all fusion constructs increase range of motion (ROM) at adjacent levels, consistent with prior biomechanical and clinical evidence [ 6 – 27 ]. However, this increase is substantially greater in two-level fusion, with the upper adjacent segment (L2–L3) showing a disproportionately larger relative increase in ROM compared to the lower adjacent segment (L5–S1), even though the latter exhibits higher absolute motion.This pattern closely mirrors the findings of Okuda et al [ 6 ], in a large longitudinal cohort, who reported significantly higher rates of upper-level degeneration. Mechanistically, fusion creates a rigid segment that shifts compensatory motion to adjacent mobile levels. In osteoporotic spines, where disc and ligamentous stiffness are already diminished [ 32 ], this compensatory hypermobility becomes even more pronounced. Beyond kinematics, internal tissue stresses offer deeper insight into degenerative mechanisms. Our research has revealed that all surgical constructs significantly elevated von Mises stresses in both the intervertebral disc (IVD) and endplate of adjacent segments, with the two-level BPS configuration generating the highest values. Crucially, the relative stress increase at L2–L3 consistently exceeded that at L5–S1, despite the greater absolute stress magnitudes observed at the caudal level—a finding consistent with the clinical reported by Okuda et al [ 6 ]. This discrepancy may stem from the smaller cross-sectional area and lower baseline load tolerance of upper lumbar discs [ 33 ]. The observed stress gradients of adjacent vertebrae - independent cage < CBT < BPS - reflect the continuous variation in structural stiffness, which is consistent with the previous results from finite element studies. [ 34 ]. Intriguingly, CBT fixation has demonstrated a moderate anterior column load, suggesting that CBT may have a potential advantage over BPS fixation in reducing intervertebral disc pressure. This potential benefit has been recently confirmed in the treatment of osteoporotic patients [ 12 , 14 ]. However, this anterior benefit comes with a trade-off in posterior loading. Under axial rotation, CBT constructs consistently generated higher facet joint stresses than BPS, particularly at the lower adjacent segment (L5–S1). During left axial rotation, the abnormally high facet joint stress is likely attributable to inherent minor asymmetries in specimen geometry, as well as variations in vertebral alignment or loading conditions. A similar biomechanical phenomenon has been reported in prior studies, particularly involving the posterior spinal structures[ 35 , 36 ]. Although the shorter length and divergent trajectory of cortical bone trajectory (CBT) screws enhance pullout resistance in osteoporotic bone, these features may concurrently reduce stability against rotational motion.[ 37 ]. As a result, greater torsional energy is transmitted through the posterior column and concentrates at the posterior facet joints. This mechanism is supported by biomechanical studies from Wang et al, [ 38 ] and Kim et al. [ 33 ], who identified facet joint stress as a key predictor of osteoarthritic progression post-fusion. Taken together, three key conclusions emerge. First, spinal fusion consistently elevates mechanical demand on adjacent segments, with two-level constructs imposing a substantially greater cumulative load than single-level procedures. Second, this burden is particularly pronounced at the upper adjacent levels, which demonstrate heightened vulnerability, as evidenced by significant increases in both range of motion and mechanical stress. This pattern is consistently supported by prior dynamic simulations and clinical observations. [ 6 , 27 ]. Third, even within fusion constructs, the use of cortical bone trajectory (CBT) fixation involves a distinct biomechanical trade-off. It reduces loading on the anterior column, which may mitigate intervertebral disc stress, but simultaneously increases stress on the posterior facet joints, particularly during rotational motion. CBT fixation appears suitable for low-demand elderly patients undergoing single-level lumbar fusion [ 17 – 20 ]. However, its use in two-level constructs or in individuals subjected to high torsional loads, such as manual laborers, should be approached with caution because of the potential for accelerated facet joint mediated degeneration[ 33 , 38 ]. It must be emphasized that ASD is not solely a biomechanical phenomenon. Systemic factors such as advanced age, high BMI, smoking, and hypertension may accelerate disc degeneration, reduce bone quality, or impair microcirculation, thereby promoting ASD [ 3 , 5 , 38 – 42 ]. These factors often act synergistically with postoperative biomechanical alterations. Our study found that cortical bone trajectory (CBT) screws provide slightly better stress distribution in the anterior column compared to traditional bicortical pedicle screws (BPS), but significantly increase stress on the facet joints of the posterior column. In osteoporotic patients, while CBT may improve screw fixation, it could impose greater mechanical load on posterior structures, necessitating careful consideration of its long-term biomechanical risks[ 43 , 44 ]. This study has several limitations. First, the model employs static, linear elastic assumptions and does not account for dynamic muscle forces or long-term biological adaptations such as bone remodeling. Second, the screw–bone interface was modeled as fully bonded, which may overestimate initial stability in osteoporotic vertebrae. Third, results are derived from a single anatomical model; inter-individual variations in facet orientation or pelvic incidence could alter load distribution. Finally, our model assumed successful fusion; however, pseudoarthrosis, which occurs relatively frequently in osteoporotic patients[ 45 , 46 ], would likely further increase the mechanical load on adjacent segments.[ 47 , 48 ]. Nevertheless, our study has been validated and shows good agreement with several previous studies, supporting its clinical relevance. A major strength of this work is its focus on two-level fusion with multiple internal fixation techniques and their effects on adjacent segments. Research on the biomechanical characteristics of adjacent segments in two-level constructs with different supplemental fixation strategies remains limited. Our findings may therefore serve as a useful reference for future studies on multi-segment adjacent segment degeneration and degenerative lumbar scoliosis. Conclusions Spinal fusion consistently increases mechanical loading on adjacent segments, with two-level constructs imposing a substantially greater cumulative load than single-level fusions. This elevated loading is particularly pronounced at the upper adjacent levels, manifesting as significant increases in both range of motion and mechanical stress. Cortical bone trajectory (CBT) screw fixation exhibits distinct biomechanical characteristics compared with traditional pedicle screw (BPS) fixation. Specifically, CBT fixation reduces stress on the anterior column but concurrently increases stress on the posterior facet joints, an effect that is especially evident during rotational motion. Abbreviations OLIF Oblique lumbar interbody fusion ASD Adjacent segment degeneration FE Finite element ROM Range of motion LIF Lumbar interbody fusion BPS bilateral pedicle screws CBT cortical bone trajectory screws Intact intact osteoporosis model L3-L4 cage L3-L4 cage model L3-L4 + BPS L3-L4 cage with bilateral pedicle screws model L3-L4 + CBT L3-L4 cage with bilateral cortical bone trajectory screws model L4-L5 cage L4-L5 cage model L4-L5 + BPS L4-L5 cage with bilateral pedicle screws model L4-L5 + CBT L4-L5 cage with bilateral cortical bone trajectory screws model L3-L5 cage L3-L5 cage model L3-L5 + BPS L3-L5 cage with bilateral pedicle screws model L3-L5 + CBT L3-L5 cage with bilateral cortical bone trajectory screws model. Declarations Ethics approval and consent to participate The present study was approved by the Ethics Committee of the Third Hospital of Hebei Medical University. Informed consent obtained from each participant was written. All protocols are carried out in accordance with relevant guidelines and regulations. Consent for publication Not applicable. Competing interests The authors declare that they have no conflict of interest. Competing interests The authors declare that they have no competing interests. Funding This work was supported by the Natural Science Foundation of Hebei Province (H2022206429). Author Contribution Conceptualization: Xianzhong Meng, Shuo Li, Lifei Wang, Hengrui ChangFormal Analysis: Xianzhong Meng, Shuo Li, Hengrui Chang, Kaibin Fan, Lifei Wang, Investigation: Jianhua Ren, Junkai Kou, Di ZhangMethodology: Shuo Li, Hengrui ChangProject Administration: Xianzhong MengWriting – Original Draft: Shuo LiWriting – Review & Editing: Shuo Li, Xianzhong Meng Acknowledgement We would like to thank all the staff in Department of Spine Surgery, the Third Hospital of Hebei Medical University for their contribution on our research. 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00:10:29","extension":"png","order_by":29,"title":"","display":"","copyAsset":false,"role":"acdc-reference","size":363112,"visible":true,"origin":"","legend":"","description":"","filename":"OnlineFIG9.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/ca479c0b8a04604aa4d058d1.png"},{"id":99187440,"identity":"36a44a02-4174-4edd-8c6e-b36a0419714d","added_by":"auto","created_at":"2025-12-30 00:10:30","extension":"xml","order_by":30,"title":"","display":"","copyAsset":false,"role":"acdc-reference","size":129347,"visible":true,"origin":"","legend":"","description":"","filename":"531a43034b7447999012100df765d80b1structuring.xml","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/9c0be7d540add1c6f558bbb9.xml"},{"id":99187438,"identity":"37770ba3-c659-460b-a721-d21f5eafbffb","added_by":"auto","created_at":"2025-12-30 00:10:30","extension":"html","order_by":31,"title":"","display":"","copyAsset":false,"role":"acdc-reference","size":138604,"visible":true,"origin":"","legend":"","description":"","filename":"earlyproof.html","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/f2bb9f59d3db36a1a6072a6c.html"},{"id":99187416,"identity":"f3abeb34-35ad-41fd-8ff5-71457a88e8a7","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":1,"title":"Figure 1","display":"","copyAsset":false,"role":"figure","size":14716208,"visible":true,"origin":"","legend":"\u003cp\u003eL3-L4 finite element models: (A)Anteroposterior view of L3-L4 cage model（B）Lateral view of L3-L4 cage model, (C) Anteroposterior view of L3-L4 cage with bilateral pedicle screws(BPS) model, (D) Latera view of L3-L4 cage with BPS model, (E) Anteroposterior view of L3-L4 cage with bilateral cortical bone trajectory screws(CBT) model, (F) Latera view of L3-L4 cage with bilateral CBT model.\u003c/p\u003e","description":"","filename":"FIG1.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/04e02357b2315d6b5d4bacf1.png"},{"id":99187426,"identity":"b370b6ab-26ef-42b0-a5ad-b561be2cf8b1","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":2,"title":"Figure 2","display":"","copyAsset":false,"role":"figure","size":14796986,"visible":true,"origin":"","legend":"\u003cp\u003eL4-L5 finite element models: (A)Anteroposterior view of L4-L5 cage model（B）Lateral view of L4-L5 cage model, (C) Anteroposterior view of L3-L4 cage with BPS model, (D) Latera view of L3-L4 cage with BPS model, (E) Anteroposterior view of L3-L4 cage with bilateral CBT model, (F) Latera view of L3-L4 cage with bilateral CBT model.\u003c/p\u003e","description":"","filename":"FIG2.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/c8fa9985b275451e9b9e80f5.png"},{"id":99187415,"identity":"74198d78-9933-4717-a5b7-26b6328f73e5","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":3,"title":"Figure 3","display":"","copyAsset":false,"role":"figure","size":14459314,"visible":true,"origin":"","legend":"\u003cp\u003eL3-L5 finite element models: (A)Anteroposterior view of L3-L5 cage model（B）Lateral view of L3-L5 cage model, (C)Anteroposterior view of L3-L5 cages BPS model, (D) Lateral view of L3-L5 cages with BPS model, (E) Anteroposterior view of L3-L5 cages with bilateral CBT model, and (F) Lateral view of L3-L5 cages with bilateral CBT model.\u003c/p\u003e","description":"","filename":"FIG3.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/23a5c7820c63a6280c5aee82.png"},{"id":99187414,"identity":"9b27a73c-f385-47d0-9449-4d1f11714de1","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":4,"title":"Figure 4","display":"","copyAsset":false,"role":"figure","size":3424993,"visible":true,"origin":"","legend":"\u003cp\u003eComparison of the range of motion (ROM) at adjacent segments and the percentage increase of ROM (%) of different models relative to the intact osteoporosis model group across different surgical constructs and loading modes. (A) The ROM of (L2-L3 segments); (B) The ROM of ( L5-S1 segments). (C) The percentage increase of ROM (%) vs. Intact at the L2-L3 level; (D) The percentage increase of ROM (%) vs. Intacl at the L5-S1 level;\u003c/p\u003e","description":"","filename":"FIG4.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/e0ef7a27aa4cb6b69d7343ce.png"},{"id":99317078,"identity":"0d38f25b-f516-4283-98ed-af4fa5337112","added_by":"auto","created_at":"2025-12-31 16:29:38","extension":"png","order_by":5,"title":"Figure 5","display":"","copyAsset":false,"role":"figure","size":1914165,"visible":true,"origin":"","legend":"\u003cp\u003eComparison of the Peak von Mises Stress in the Intervertebral Discs(IVD) at adjacent segments and the percentage increase of DISC stress of different models relative to the intact osteoporosis model group across different surgical constructs and loading modes. (A) The IVD Stress of (L2-L3 segments); (B) The IVD Stress of ( L5-S1 segments). (C) L2-L3 Intervertebral Disc Stress Increase (%) vs. Intact; (D) L5-S1 Intervertebral Disc Stress Increase (%) vs. Intact\u003c/p\u003e","description":"","filename":"FIG5.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/369372f071689b71750eadb3.png"},{"id":99187433,"identity":"dc04ffe0-e4ef-40be-a51c-e061938aa555","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":6,"title":"Figure 6","display":"","copyAsset":false,"role":"figure","size":2654254,"visible":true,"origin":"","legend":"\u003cp\u003eComparison of the Endplate Stress in the Intervertebral Discs(IVD) at adjacent segments and the percentage increase of Endplate stress of different models relative to the intact osteoporosis model group across different surgical constructs and loading modes. (A) The Endplate Stress of (L2-L3 segments); (B) The Endplate Stres of ( L5-S1 segments). (C) L2-L3 Endplate Stress Increase (%) vs. Intact; (D) L5-S1Endplate Stress Increase (%) vs. Intact\u003c/p\u003e","description":"","filename":"FIG6.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/5abff83ddaa7c6c6184fac7b.png"},{"id":99187434,"identity":"8612728a-aba6-4ad6-8f12-918e8d8c9119","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":7,"title":"Figure 7","display":"","copyAsset":false,"role":"figure","size":2361379,"visible":true,"origin":"","legend":"\u003cp\u003eComparison of the Facet Joint Stress at adjacent segments and the percentage increase of Facet Joint stress of different models relative to the intact osteoporosis model group across different surgical constructs and loading modes. (A) The Facet Joint Stress of (L2-L3 segments); (B) The Facet Joint Stres of ( L5-S1 segments). (C) L2-L3 Facet Joint Stress Increase (%) vs. Intact; (D) L5-S1 Facet Joint Stress Increase (%) vs. Intact\u003c/p\u003e","description":"","filename":"FIG7.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/265047da86896405b4b31405.png"},{"id":99187424,"identity":"c1a5e7a4-bc72-45a5-84f2-e11bb8f46186","added_by":"auto","created_at":"2025-12-30 00:10:29","extension":"png","order_by":8,"title":"Figure 8","display":"","copyAsset":false,"role":"figure","size":11484826,"visible":true,"origin":"","legend":"\u003cp\u003ePeak stress distribution in the L2–L3 and L5–S1 intervertebral discs across different surgical models. Pictures of each group of models from top to bottom: L2–L3 cage model, L3–L5 cage model, L2–L3 + BPS model, L3–L5 + BPS model, L2–L3 + CBT model, and L3–L5 + CBT model. (A) shows the stress distribution in the L2–L3 intervertebral disc; (B) shows the peak stress distribution in the L5–S1 intervertebral disc.\u003c/p\u003e","description":"","filename":"FIG8.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/a7e8086a4c428e3f87ca2d2e.png"},{"id":99187442,"identity":"4cade268-a465-4b6b-973f-a0a02f20df5a","added_by":"auto","created_at":"2025-12-30 00:10:30","extension":"png","order_by":9,"title":"Figure 9","display":"","copyAsset":false,"role":"figure","size":12721160,"visible":true,"origin":"","legend":"\u003cp\u003ePeak stress distribution in the L2–L3 and L5–S1 endplate across different surgical models. Pictures of each group of models from top to bottom: L2–L3 cage model, L3–L5 cage model, L2–L3 + BPS model, L3–L5 + BPS model, L2–L3 + CBT model, and L3–L5 + CBT model. (A) shows the stress distribution in the L2–L3 endplate; (B) shows the peak stress distribution in the L5–S1 endplate.\u003c/p\u003e","description":"","filename":"FIG9.png","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/80a66a87e1a68075142d4764.png"},{"id":99787944,"identity":"5208268a-7949-4a25-864f-317673bc1908","added_by":"auto","created_at":"2026-01-08 12:42:05","extension":"pdf","order_by":0,"title":"","display":"","copyAsset":false,"role":"manuscript-pdf","size":72960429,"visible":true,"origin":"","legend":"","description":"","filename":"manuscript.pdf","url":"https://assets-eu.researchsquare.com/files/rs-8272176/v1/623c8ee1-773f-403b-a6ed-48314ce3b205.pdf"}],"financialInterests":"No competing interests reported.","formattedTitle":"Biomechanical Changes of Adjacent Segments Following Single- versus Two-level Oblique Lateral Interbody Fusion with Different Fixation Methods: A Finite Element study","fulltext":[{"header":"Background","content":"\u003cp\u003eLumbar spinal fusion is widely employed for the treatment of degenerative and unstable spinal disorders. However, complications including implant failure, cage subsidence, and adjacent segment degeneration (ASD) remain prevalent, particularly in elderly patients with compromised bone quality. Osteoporosis, which is common in this group, weakens bone and alters spinal biomechanics, raising the risk of screw loosening and construct instability [\u003cspan citationid=\"CR1\" class=\"CitationRef\"\u003e1\u003c/span\u003e, \u003cspan citationid=\"CR2\" class=\"CitationRef\"\u003e2\u003c/span\u003e]. ASD is a major reason for reoperation and long-term disability after lumbar fusion, with 10-year incidence rates reported between 15% and 18% [\u003cspan additionalcitationids=\"CR4 CR5\" citationid=\"CR3\" class=\"CitationRef\"\u003e3\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eOLIF has gained widespread adoption in the management of degenerative lumbar disorders due to its capacity to restore disc height, correct sagittal alignment, and achieve high fusion rates with minimal soft tissue disruption[\u003cspan additionalcitationids=\"CR8\" citationid=\"CR7\" class=\"CitationRef\"\u003e7\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR9\" class=\"CitationRef\"\u003e9\u003c/span\u003e]. While stand alone OLIF may provide adequate stability in select single level pathologies, multilevel constructs often require supplemental posterior fixation to mitigate the risk of cage subsidence and pseudarthrosis, especially in osteoporotic patients[\u003cspan citationid=\"CR10\" class=\"CitationRef\"\u003e10\u003c/span\u003e, \u003cspan citationid=\"CR11\" class=\"CitationRef\"\u003e11\u003c/span\u003e]. Compared with the standard fixation method of bilateral pedicle screws (BPS) for lumbar fusion, cortical bone trajectory (CBT) screws have been reported in recent years to provide similar pull-out strength and resistance to buckling in osteoporotic bone, with less soft tissue injury and faster recovery[\u003cspan additionalcitationids=\"CR13 CR14 CR15\" citationid=\"CR12\" class=\"CitationRef\"\u003e12\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR16\" class=\"CitationRef\"\u003e16\u003c/span\u003e]. Moreover, some clinical data have shown that cortical bone trajectory (CBT) screws and bilateral pedicle screws (BPS) have similar fusion rates and patient-reported outcomes in single-level surgery[\u003cspan additionalcitationids=\"CR18 CR19\" citationid=\"CR17\" class=\"CitationRef\"\u003e17\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR20\" class=\"CitationRef\"\u003e20\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eCurrently, there are various studies on the impact of single-level OLIF surgery on adjacent segment vertebrae, including clinical studies, in vitro biomechanical studies, and finite element studies[\u003cspan additionalcitationids=\"CR22 CR23\" citationid=\"CR21\" class=\"CitationRef\"\u003e21\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR24\" class=\"CitationRef\"\u003e24\u003c/span\u003e]. To date, no study has systematically investigated how these techniques influence load distribution at both the superior and inferior adjacent segments across different fusion configurations, including single-level and two-level fusions, under physiologically relevant loading conditions.To address this, we developed and validated a detailed finite element model of the osteoporotic L1-S1 lumbar spine. This study aimed to investigate the biomechanical effects of single- and two-level OLIF with different internal fixation methods on adjacent segments in an osteoporotic lumbar spine model.\u003c/p\u003e"},{"header":"Methods","content":"\u003cdiv id=\"Sec3\" class=\"Section2\"\u003e \u003ch2\u003eConstruction of an Intact Lumbar Finite Element Model\u003c/h2\u003e \u003cp\u003eThe finite element (FE) modeling approach used in this study is identical to that established by Fan et al[\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e]. from our research group for biomechanical analysis of osteoporotic lumbar spines following OLIF. Their study investigated the biomechanics of two-level OLIF in osteoporotic spines with different internal fixation techniques and found that supplemental fixation enhanced segmental stability and reduced cage stress. BPS outperformed unilateral pedicle screws (UPS) and CBT screws in restricting segmental motion and lowering stresses on cages and implants. Here, we applied the same FE model to evaluate the biomechanical effects of single- and two-level fixation combined with OLIF on adjacent segments in an osteoporotic spine. Briefly, a three-dimensional FE model of the intact L1-S1 lumbar spine was reconstructed from computed tomography (CT) scans of a 30-year-old healthy male volunteer (slice thickness: 0.625 mm; 570 slices in DICOM format). Segmentation and initial 3D geometry reconstruction were performed using Mimics 21.0 (Materialise, Leuven, Belgium), followed by surface refinement in Geomagic Studio 12.0 (3D Systems, Rock Hill, SC, USA). Intervertebral discs and other soft tissues were modeled in Creo 8.0 (PTC, Boston, MA, USA), and the final anatomical structures were meshed using ANSA (BETA CAE Systems, Thessaloniki, Greece).\u003c/p\u003e \u003cp\u003eThe model includes vertebral bodies (L1-S1), posterior bony elements (pedicles, laminae, facets), cartilaginous endplates, intervertebral discs (annulus fibrosus and nucleus pulposus), and seven major spinal ligaments (anterior longitudinal ligaments, posterior longitudinal ligaments, ligamentum flavum, supraspinous ligaments, interspinous ligaments, capsular ligaments, and intertransverse ligaments). A mixed mesh of tetrahedral and hexahedral elements was used, with an element size of 1 mm determined through convergence analysis (ROM variation\u0026thinsp;\u0026lt;\u0026thinsp;1% between 1 mm and 2 mm meshes under 400 N compression\u0026thinsp;+\u0026thinsp;7.5 Nm flexion).\u003c/p\u003e \u003cp\u003eMaterial properties were assigned based on established literature [\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e]: cortical bone (E\u0026thinsp;=\u0026thinsp;12.0 GPa), cancellous bone (E\u0026thinsp;=\u0026thinsp;100 MPa), endplates (E\u0026thinsp;=\u0026thinsp;4 GPa), and posterior elements (E\u0026thinsp;=\u0026thinsp;3.5 GPa). To simulate osteoporosis, the elastic modulus of cortical, the endplates, and the posterior elements were reduced by 33%, and the cancellous bone were reduced 66% respectively [\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e, \u003cspan citationid=\"CR26\" class=\"CitationRef\"\u003e26\u003c/span\u003e]. The annulus fibrosus was modeled as a fiber-reinforced composite with 12 layers of collagen fibers oriented at \u0026plusmn;\u0026thinsp;30\u0026deg;(E\u0026thinsp;=\u0026thinsp;4.2 MPa), while the nucleus pulposus was treated as nearly incompressible (E\u0026thinsp;=\u0026thinsp;0.1 MPa, ν\u0026thinsp;=\u0026thinsp;0.49). Ligaments were represented as nonlinear tension-only truss elements, and facet joints were defined as surface-to-surface contact pairs with a friction coefficient of 0.1. The detailed material properties of each component are listed in Table\u0026nbsp;\u003cspan refid=\"Tab1\" class=\"InternalRef\"\u003e1\u003c/span\u003e.\u003c/p\u003e \u003cp\u003e \u003cdiv class=\"gridtable\"\u003e\u003ctable float=\"Yes\" id=\"Tab1\" border=\"1\"\u003e \u003ccaption language=\"En\"\u003e \u003cdiv class=\"CaptionNumber\"\u003eTable 1\u003c/div\u003e \u003cdiv class=\"CaptionContent\"\u003e \u003cp\u003eMaterial properties of the FEM and implants\u003c/p\u003e \u003c/div\u003e \u003c/caption\u003e \u003ccolgroup cols=\"3\"\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c1\" colnum=\"1\"\u003e\u003c/div\u003e \u003cdiv align=\"left\" class=\"colspec\" colname=\"c2\" colnum=\"2\"\u003e\u003c/div\u003e \u003cdiv align=\"char\" char=\".\" class=\"colspec\" colname=\"c3\" colnum=\"3\"\u003e\u003c/div\u003e \u003cthead\u003e \u003ctr\u003e \u003cth align=\"left\" colname=\"c1\"\u003e \u003cp\u003eComponents\u003c/p\u003e \u003c/th\u003e \u003cth align=\"left\" colname=\"c2\"\u003e \u003cp\u003eYoung\u0026rsquo;s modulus (MPa)\u003c/p\u003e \u003c/th\u003e \u003cth align=\"left\" colname=\"c3\"\u003e \u003cp\u003ePoisson ratio\u003c/p\u003e \u003c/th\u003e \u003c/tr\u003e \u003c/thead\u003e \u003ctbody\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eCortical bone\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e8040 (\u0026ldquo;normal\u0026rdquo;: 12000)\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eCancellous bone\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e34 (\u0026ldquo;normal\u0026rdquo;: 100)\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.2\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eEnd-plate\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e2640 (\u0026ldquo;normal\u0026rdquo;: 4000)\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.4\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003ePosterior elements\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e2345 (\u0026ldquo;normal\u0026rdquo;: 3500)\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.25\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eNucleus pulposus\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e1\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.49\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eAnnulus\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e4.2\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.45\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eAnterior longitudinal ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e20\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003ePosterior longitudinal ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e20\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eLigamentum flavum\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e19.5\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eInterspinous ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e11.6\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eSupraspinous ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e15\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eIntertransverse ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e58.7\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eCapsular ligament\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e32.9\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eCage\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e3500\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003ctr\u003e \u003ctd align=\"left\" colname=\"c1\"\u003e \u003cp\u003eScrew and rod\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"left\" colname=\"c2\"\u003e \u003cp\u003e110,000\u003c/p\u003e \u003c/td\u003e \u003ctd align=\"char\" char=\".\" colname=\"c3\"\u003e \u003cp\u003e0.3\u003c/p\u003e \u003c/td\u003e \u003c/tr\u003e \u003c/tbody\u003e \u003c/colgroup\u003e \u003c/table\u003e\u003c/div\u003e \u003c/p\u003e \u003c/div\u003e\n\u003ch3\u003eBoundary and Loading Conditions\u003c/h3\u003e\n\u003cp\u003eBoundary and loading conditions followed standard protocols: the inferior endplate of S1 was fully fixed (all six degrees of freedom constrained), a 400 N follower load was applied to the superior endplate of L1, and a pure moment of 10 Nm was subsequently applied in six directions (flexion, extension, left bending, right bending, and left rotation, and right rotation) to simulate physiological spinal motion.\u003c/p\u003e\n\u003ch3\u003eSurgical Models and Fixation Configurations\u003c/h3\u003e\n\u003cp\u003eThree-dimensional geometric models of the internal fixation devices were developed in Creo 8.0 using the Part module, based on the actual dimensions of the interbody cage and supplemental fixation components. The interbody cage was modeled after the Oracle cage (DePuy Synthes), measuring 40 mm in length, 22 mm in width, 11 mm in anterior height, 8 mm in posterior height, and featuring an 8\u0026deg; lordotic angle. The pedicle screws had a diameter of 6.5 mm and a length of 50 mm, while the cortical bone screws were 5.0 mm in diameter and 30 mm long. The connecting rods had a diameter of 5.5 mm. Both the cage and the supplemental fixation devices were discretized using tetrahedral meshing. Bone-screw interfaces were assumed to be fully bonded to simulate immediate postoperative stability. Detailed material properties for each component are provided in Table\u0026nbsp;\u003cspan refid=\"Tab1\" class=\"InternalRef\"\u003e1\u003c/span\u003e.\u003c/p\u003e \u003cp\u003eThe surgical segment was defined as the L3\u0026ndash;L5 intervertebral space, and the annulus fibrosus, nucleus pulposus, and cartilaginous endplate were removed from the left side. Nine surgical models were created to evaluate the biomechanical performance of different fusion strategies under osteoporotic conditions in Fig.\u0026nbsp;\u003cspan refid=\"Fig1\" class=\"InternalRef\"\u003e1\u003c/span\u003e\u0026ndash;\u003cspan refid=\"Fig3\" class=\"InternalRef\"\u003e3\u003c/span\u003e:\u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003cp\u003eSingle-level L3-L4 fusion: L3-4 Cage: Standalone polyetheretherketone (PEEK) interbody cage; L3-4 BPS: Cage\u0026thinsp;+\u0026thinsp;bilateral pedicle screw system (6.5 mm diameter, 50 mm length); L3-4 CBT: Cage\u0026thinsp;+\u0026thinsp;bilateral cortical bone trajectory (CBT) screws (5.0 mm diameter, 30 mm length, 10\u0026deg; medial, 25\u0026deg; cephalad). Single-level L4-L5 fusion: L4-5 Cage, L4-5\u0026thinsp;+\u0026thinsp;BPS, L4-5 CBT (same screw parameters)\u003c/p\u003e \u003cp\u003eTwo-level L3-L5 fusion: L3-5 Cage: Cages at both levels without posterior fixation; L3-5 BPS: Cages\u0026thinsp;+\u0026thinsp;bilateral pedicle screws at L3 and L5; L3-5 CBT: Cages\u0026thinsp;+\u0026thinsp;bilateral CBT screws at L3 and L5.\u003c/p\u003e \u003cp\u003eImportantly, while the FE model and surgical configurations (including fusion levels, cage types, and fixation strategies) are identical to those in Fan et al. [\u003cspan citationid=\"CR20\" class=\"CitationRef\"\u003e20\u003c/span\u003e], the focus of the present study differs substantially. Whereas Fan et al. primarily investigated the biomechanical behavior of the instrumented (surgical) segments and internal fixation devices, this study specifically evaluates the biomechanical response of the adjacent non-fused segments, the cranial adjacent segment (L2-L3) and caudal adjacent segment ( L5-S1).Our research emphasis on range of motion (ROM) of adjacent vertebral segments, intradiscal stress distribution, and facet joint contact forces. These parameters are critical indicators for assessing the potential risk of adjacent segment degeneration (ASD) following lumbar fusion under osteoporotic conditions.\u003c/p\u003e"},{"header":"Results","content":"\u003cdiv id=\"Sec7\" class=\"Section2\"\u003e \u003ch2\u003eRange of Motion (ROM)\u003c/h2\u003e \u003cp\u003eAt the L2-L3 level, all instrumented models, demonstrated increased ROM relative to IntactOP model (flexion: 5.02\u0026deg;), confirming compensatory hypermobility due to increased caudal stiffness. The largest increases were observed in flexion (up to 6.70\u0026deg; in L3-5cage\u0026thinsp;+\u0026thinsp;BPS; +33.5%) and lateral bending (right bend: 6.26\u0026deg; vs. 5.23\u0026deg;; +19.7%). Similarly, at L5-S1, ROM consistently exceeded baseline values (IntactOP model flexion: 11.39\u0026deg;), with the greatest elevation seen in the L3-5cage\u0026thinsp;+\u0026thinsp;BPS model (flexion: 14.69\u0026deg;; +29.0%). These findings indicate that both proximal and distal adjacent segments experience significant kinematic overload following lumbar fusion.\u003c/p\u003e \u003cp\u003eTwo-level fusion (L3-5) consistently induced greater ROM increases at adjacent segments than single-level fusion (L3-4 or L4-5). At L2-L3, flexion ROM rose from 5.57\u0026deg; (L3-4cage) to 6.02\u0026deg; (L3-5cage; +8.1%), and further to 6.70\u0026deg;with BPS fixation (L3-5cage\u0026thinsp;+\u0026thinsp;BPS; +12.2% vs. L3-4cage\u0026thinsp;+\u0026thinsp;BPS). At L5-S1, L3-5 fusion models exhibited markedly higher motion than L4-5 counterparts: for example, flexion ROM was 14.69\u0026deg; (L3-5cage\u0026thinsp;+\u0026thinsp;BPS) versus 13.14\u0026deg; (L4-5cage\u0026thinsp;+\u0026thinsp;BPS; +11.8%). This trend held across all loading modes. Two-level fusion (L3-L5) resulted in even greater compensatory motion at both adjacent levels: L2-L3 ROM increased by 19.98% to 33.33%, while L5-S1 ROM increased by 15.20% to 28.99%, confirming that extending the fusion span amplifies adjacent segmental motion. Notably, the greatest ROM increases were observed in the BPS groups across all fusion types, whereas CBT fixation consistently demonstrated lower adjacent segment mobility compared to BPS. Figure\u0026nbsp;\u003cspan refid=\"Fig4\" class=\"InternalRef\"\u003e4\u003c/span\u003e.\u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003c/div\u003e \u003cdiv id=\"Sec8\" class=\"Section2\"\u003e \u003ch2\u003eIntervertebral Disc Stress\u003c/h2\u003e \u003cp\u003ePeak stresses in the intervertebral discs were elevated in all fusion models. Single-level fusion consistently elevated IVD stress compared to the intact condition, with posterior instrumentation exacerbating this effect. At the proximal adjacent segment (L2\u0026ndash;L3), the L3\u0026ndash;L4 cage-only construct increased disc stress to 1.79 MPa in flexion (a 16.2% increase from intact 1.54 MPa). The addition of bilateral pedicle screws (BPS) further raised stress to 2.02 MPa (+\u0026thinsp;31.2%), while cortical bone trajectory (CBT) screws yielded an intermediate value of 1.95 MPa (+\u0026thinsp;26.6%). A similar hierarchy was observed at the distal adjacent segment (L5\u0026ndash;S1): L4\u0026ndash;L5 cage increased stress to 3.06 MPa in flexion (intact: 2.75 MPa; +11.3%), whereas L4\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS and +\u0026thinsp;CBT reached 3.34 MPa (+\u0026thinsp;21.5%) and 3.24 MPa (+\u0026thinsp;18.2%), respectively. These trends were consistent across all loading modes, indicating that BPS fixation imposes the greatest additional load on adjacent discs, followed by CBT and cage-alone constructs.\u003c/p\u003e \u003cp\u003eExtending fusion to two levels (L3\u0026ndash;L5) further amplified adjacent disc stress beyond single-level scenarios. For example, at L2\u0026ndash;L3, the L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS model produced a peak stress of 2.20 MPa in flexion\u0026mdash;18% higher than the 1.88 MPa observed with L3\u0026ndash;L5 cage and 9% higher than L3\u0026ndash;L4\u0026thinsp;+\u0026thinsp;BPS (2.02 MPa). Similarly, at L5\u0026ndash;S1, L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS generated 3.67 MPa in flexion, exceeding both L3\u0026ndash;L5 cage (3.32 MPa) and L4\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS (3.34 MPa). Across all constructs, two-level fusion increased adjacent IVD stress by approximately 10\u0026ndash;13 percentage points compared to anatomically matched single-level fusions, underscoring the cumulative biomechanical burden of longer constructs.\u003c/p\u003e \u003cp\u003eNotably, the proximal adjacent segment (L2\u0026ndash;L3) exhibited greater relative stress elevation than the distal segment (L5\u0026ndash;S1) under identical fusion conditions. In the L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS model, L2\u0026ndash;L3 disc stress increased by 42.9% in flexion (from 1.54 to 2.20 MPa), whereas L5\u0026ndash;S1 stress rose by 33.5% (from 2.75 to 3.67 MPa)\u0026mdash;a difference of 9.4 percentage points. This pattern persisted across all two-level constructs and loading directions, with L2\u0026ndash;L3 consistently demonstrating 7\u0026ndash;12% higher relative stress increases than L5\u0026ndash;S1. Despite the higher baseline stress in the intact L5\u0026ndash;S1 disc (e.g., 2.75 MPa vs. 1.54 MPa in flexion), the proportional mechanical perturbation was more pronounced at the cranial adjacent level. (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003e). The peak stress distribution in the L2\u0026ndash;L3 and L5\u0026ndash;S1 intervertebral discs under single-level and two-level surgical configurations as shown in Fig.\u0026nbsp;\u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e8\u003c/span\u003e.\u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003c/div\u003e\n\u003ch3\u003eEndplate Stress\u003c/h3\u003e\n\u003cp\u003eEndplate stress followed a pattern closely aligned with disc stress. At L2-L3, peak endplate stress increased by 20\u0026ndash;52% in flexion and 19\u0026ndash;45% in extension, reaching 13.82 MPa (L3-L5\u0026thinsp;+\u0026thinsp;BPS) in flexion and 10.34 MPa in extension. At L5-S1, stress rose by 14\u0026ndash;43% in flexion and 13\u0026ndash;35% in extension, with the L3-L5\u0026thinsp;+\u0026thinsp;BPS model again showing the highest values (16.28 MPa in flexion; 11.28 MPa in extension). The L3-L5\u0026thinsp;+\u0026thinsp;CBT construct also induced substantial endplate loading (15.76 MPa in flexion), approaching that of BPS.\u003c/p\u003e \u003cp\u003eStand-alone cage fusion consistently demonstrated the smallest stress increases (13\u0026ndash;20%) across all critical loading modes, significantly lower than constructs augmented with BPS or CBT screws (25\u0026ndash;52% increase). Although CBT fixation yielded slightly lower stresses than BPS, it did not confer a substantial biomechanical advantage in mitigating adjacent segment loading. (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e6\u003c/span\u003e) The peak stress distribution in the L2-L3 and L5-S1 endplate under single-level and two-level surgical configurations as shown in Fig.\u0026nbsp;\u003cspan refid=\"Fig8\" class=\"InternalRef\"\u003e9\u003c/span\u003e.\u003c/p\u003e \u003cp\u003e \u003c/p\u003e \u003cp\u003e \u003c/p\u003e\n\u003ch3\u003eFacet Joint Stress\u003c/h3\u003e\n\u003cp\u003eIn stark contrast to the anterior column, facet joint stresses were predominantly elevated during extension and rotation. At L2-L3, extension stress surged from 16.85 MPa (intact) to 26.02 MPa in the L3-L5\u0026thinsp;+\u0026thinsp;CBT model (+\u0026thinsp;54%), the highest value observed across all conditions. During rotation, L3-L5\u0026thinsp;+\u0026thinsp;CBT also produced the greatest stresses (22.25 MPa left; 22.38 MPa right). At L5-S1, extension stress increased by 24\u0026ndash;76%, with L3-L5\u0026thinsp;+\u0026thinsp;BPS reaching 15.56 MPa (+\u0026thinsp;76%). Most strikingly, during left axial rotation, the L3-L5 CBT construct generated an exceptionally high facet stress of 17.64 MPa, representing a 143% increase compared with the intact condition (7.25 MPa) and markedly exceeding the values observed in all other models. In contrast, facet stresses remained minimal during flexion (\u0026lt;\u0026thinsp;0.25 MPa at L2-L3; \u0026lt;4.08 MPa at L5-S1), confirming their non-load-bearing role in this motion. (Fig.\u0026nbsp;\u003cspan refid=\"Fig9\" class=\"InternalRef\"\u003e7\u003c/span\u003e).\u003c/p\u003e \u003cp\u003e \u003c/p\u003e"},{"header":"Discussion","content":"\u003cp\u003eAdjacent segment degeneration (ASD) remains a leading cause of reoperation after lumbar fusion, with 10-year rates of 15% to 18% [\u003cspan citationid=\"CR3\" class=\"CitationRef\"\u003e3\u003c/span\u003e, \u003cspan citationid=\"CR5\" class=\"CitationRef\"\u003e5\u003c/span\u003e]. Biomechanical overload at adjacent levels\u0026mdash;particularly the upper segment, where degeneration is consistently more severe [\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e, \u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]\u0026mdash;is a primary driver, and this risk is amplified in osteoporotic patients due to poor bone quality and reduced screw\u0026ndash;bone stability [\u003cspan citationid=\"CR1\" class=\"CitationRef\"\u003e1\u003c/span\u003e, \u003cspan citationid=\"CR2\" class=\"CitationRef\"\u003e2\u003c/span\u003e]. While cortical bone trajectory (CBT) screws offer improved pullout strength in low-density bone compared to conventional pedicle screws [\u003cspan citationid=\"CR12\" class=\"CitationRef\"\u003e12\u003c/span\u003e, \u003cspan citationid=\"CR14\" class=\"CitationRef\"\u003e14\u003c/span\u003e], their impact on adjacent segment loading in multilevel constructs remains poorly understood. Most finite element studies have focused on single-level fusion in non-osteoporotic models [\u003cspan additionalcitationids=\"CR29\" citationid=\"CR28\" class=\"CitationRef\"\u003e28\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR30\" class=\"CitationRef\"\u003e30\u003c/span\u003e], limiting clinical relevance for elderly patients who often present with contiguous multilevel disease. To address this gap, we developed a validated osteoporotic L1\u0026ndash;S1 finite element model to compare three fixation strategies\u0026mdash;standalone interbody cage, bilateral pedicle screws (BPS), and CBT screws\u0026mdash;under both single-level (L4\u0026ndash;L5) and two-level (L3\u0026ndash;L5) fusion scenarios, with L3\u0026ndash;L5 selected as a representative multilevel degeneration pattern in older adults [\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e]. Model kinematics were validated against in vitro motion data [\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e, \u003cspan citationid=\"CR31\" class=\"CitationRef\"\u003e31\u003c/span\u003e] to ensure physiological fidelity.\u003c/p\u003e \u003cp\u003eOur kinematic analyses confirm that all fusion constructs increase range of motion (ROM) at adjacent levels, consistent with prior biomechanical and clinical evidence [\u003cspan additionalcitationids=\"CR7 CR8 CR9 CR10 CR11 CR12 CR13 CR14 CR15 CR16 CR17 CR18 CR19 CR20 CR21 CR22 CR23 CR24 CR25 CR26\" citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. However, this increase is substantially greater in two-level fusion, with the upper adjacent segment (L2\u0026ndash;L3) showing a disproportionately larger relative increase in ROM compared to the lower adjacent segment (L5\u0026ndash;S1), even though the latter exhibits higher absolute motion.This pattern closely mirrors the findings of Okuda et al [\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e], in a large longitudinal cohort, who reported significantly higher rates of upper-level degeneration. Mechanistically, fusion creates a rigid segment that shifts compensatory motion to adjacent mobile levels. In osteoporotic spines, where disc and ligamentous stiffness are already diminished [\u003cspan citationid=\"CR32\" class=\"CitationRef\"\u003e32\u003c/span\u003e], this compensatory hypermobility becomes even more pronounced.\u003c/p\u003e \u003cp\u003eBeyond kinematics, internal tissue stresses offer deeper insight into degenerative mechanisms. Our research has revealed that all surgical constructs significantly elevated von Mises stresses in both the intervertebral disc (IVD) and endplate of adjacent segments, with the two-level BPS configuration generating the highest values. Crucially, the relative stress increase at L2\u0026ndash;L3 consistently exceeded that at L5\u0026ndash;S1, despite the greater absolute stress magnitudes observed at the caudal level\u0026mdash;a finding consistent with the clinical reported by Okuda et al [\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e]. This discrepancy may stem from the smaller cross-sectional area and lower baseline load tolerance of upper lumbar discs [\u003cspan citationid=\"CR33\" class=\"CitationRef\"\u003e33\u003c/span\u003e]. The observed stress gradients of adjacent vertebrae - independent cage\u0026thinsp;\u0026lt;\u0026thinsp;CBT\u0026thinsp;\u0026lt;\u0026thinsp;BPS - reflect the continuous variation in structural stiffness, which is consistent with the previous results from finite element studies. [\u003cspan citationid=\"CR34\" class=\"CitationRef\"\u003e34\u003c/span\u003e]. Intriguingly, CBT fixation has demonstrated a moderate anterior column load, suggesting that CBT may have a potential advantage over BPS fixation in reducing intervertebral disc pressure. This potential benefit has been recently confirmed in the treatment of osteoporotic patients [\u003cspan citationid=\"CR12\" class=\"CitationRef\"\u003e12\u003c/span\u003e, \u003cspan citationid=\"CR14\" class=\"CitationRef\"\u003e14\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eHowever, this anterior benefit comes with a trade-off in posterior loading. Under axial rotation, CBT constructs consistently generated higher facet joint stresses than BPS, particularly at the lower adjacent segment (L5\u0026ndash;S1). During left axial rotation, the abnormally high facet joint stress is likely attributable to inherent minor asymmetries in specimen geometry, as well as variations in vertebral alignment or loading conditions. A similar biomechanical phenomenon has been reported in prior studies, particularly involving the posterior spinal structures[\u003cspan citationid=\"CR35\" class=\"CitationRef\"\u003e35\u003c/span\u003e, \u003cspan citationid=\"CR36\" class=\"CitationRef\"\u003e36\u003c/span\u003e]. Although the shorter length and divergent trajectory of cortical bone trajectory (CBT) screws enhance pullout resistance in osteoporotic bone, these features may concurrently reduce stability against rotational motion.[\u003cspan citationid=\"CR37\" class=\"CitationRef\"\u003e37\u003c/span\u003e]. As a result, greater torsional energy is transmitted through the posterior column and concentrates at the posterior facet joints. This mechanism is supported by biomechanical studies from Wang et al, [\u003cspan citationid=\"CR38\" class=\"CitationRef\"\u003e38\u003c/span\u003e] and Kim et al. [\u003cspan citationid=\"CR33\" class=\"CitationRef\"\u003e33\u003c/span\u003e], who identified facet joint stress as a key predictor of osteoarthritic progression post-fusion.\u003c/p\u003e \u003cp\u003eTaken together, three key conclusions emerge. First, spinal fusion consistently elevates mechanical demand on adjacent segments, with two-level constructs imposing a substantially greater cumulative load than single-level procedures. Second, this burden is particularly pronounced at the upper adjacent levels, which demonstrate heightened vulnerability, as evidenced by significant increases in both range of motion and mechanical stress. This pattern is consistently supported by prior dynamic simulations and clinical observations. [\u003cspan citationid=\"CR6\" class=\"CitationRef\"\u003e6\u003c/span\u003e, \u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. Third, even within fusion constructs, the use of cortical bone trajectory (CBT) fixation involves a distinct biomechanical trade-off. It reduces loading on the anterior column, which may mitigate intervertebral disc stress, but simultaneously increases stress on the posterior facet joints, particularly during rotational motion. CBT fixation appears suitable for low-demand elderly patients undergoing single-level lumbar fusion [\u003cspan additionalcitationids=\"CR18 CR19\" citationid=\"CR17\" class=\"CitationRef\"\u003e17\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR20\" class=\"CitationRef\"\u003e20\u003c/span\u003e]. However, its use in two-level constructs or in individuals subjected to high torsional loads, such as manual laborers, should be approached with caution because of the potential for accelerated facet joint mediated degeneration[\u003cspan citationid=\"CR33\" class=\"CitationRef\"\u003e33\u003c/span\u003e, \u003cspan citationid=\"CR38\" class=\"CitationRef\"\u003e38\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eIt must be emphasized that ASD is not solely a biomechanical phenomenon. Systemic factors such as advanced age, high BMI, smoking, and hypertension may accelerate disc degeneration, reduce bone quality, or impair microcirculation, thereby promoting ASD [\u003cspan citationid=\"CR3\" class=\"CitationRef\"\u003e3\u003c/span\u003e, \u003cspan citationid=\"CR5\" class=\"CitationRef\"\u003e5\u003c/span\u003e, \u003cspan additionalcitationids=\"CR39 CR40 CR41\" citationid=\"CR38\" class=\"CitationRef\"\u003e38\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR42\" class=\"CitationRef\"\u003e42\u003c/span\u003e]. These factors often act synergistically with postoperative biomechanical alterations. Our study found that cortical bone trajectory (CBT) screws provide slightly better stress distribution in the anterior column compared to traditional bicortical pedicle screws (BPS), but significantly increase stress on the facet joints of the posterior column. In osteoporotic patients, while CBT may improve screw fixation, it could impose greater mechanical load on posterior structures, necessitating careful consideration of its long-term biomechanical risks[\u003cspan citationid=\"CR43\" class=\"CitationRef\"\u003e43\u003c/span\u003e, \u003cspan citationid=\"CR44\" class=\"CitationRef\"\u003e44\u003c/span\u003e].\u003c/p\u003e \u003cp\u003eThis study has several limitations. First, the model employs static, linear elastic assumptions and does not account for dynamic muscle forces or long-term biological adaptations such as bone remodeling. Second, the screw\u0026ndash;bone interface was modeled as fully bonded, which may overestimate initial stability in osteoporotic vertebrae. Third, results are derived from a single anatomical model; inter-individual variations in facet orientation or pelvic incidence could alter load distribution. Finally, our model assumed successful fusion; however, pseudoarthrosis, which occurs relatively frequently in osteoporotic patients[\u003cspan citationid=\"CR45\" class=\"CitationRef\"\u003e45\u003c/span\u003e, \u003cspan citationid=\"CR46\" class=\"CitationRef\"\u003e46\u003c/span\u003e], would likely further increase the mechanical load on adjacent segments.[\u003cspan citationid=\"CR47\" class=\"CitationRef\"\u003e47\u003c/span\u003e, \u003cspan citationid=\"CR48\" class=\"CitationRef\"\u003e48\u003c/span\u003e]. Nevertheless, our study has been validated and shows good agreement with several previous studies, supporting its clinical relevance. A major strength of this work is its focus on two-level fusion with multiple internal fixation techniques and their effects on adjacent segments. Research on the biomechanical characteristics of adjacent segments in two-level constructs with different supplemental fixation strategies remains limited. Our findings may therefore serve as a useful reference for future studies on multi-segment adjacent segment degeneration and degenerative lumbar scoliosis.\u003c/p\u003e"},{"header":"Conclusions","content":"\u003cp\u003eSpinal fusion consistently increases mechanical loading on adjacent segments, with two-level constructs imposing a substantially greater cumulative load than single-level fusions. This elevated loading is particularly pronounced at the upper adjacent levels, manifesting as significant increases in both range of motion and mechanical stress.\u003c/p\u003e \u003cp\u003eCortical bone trajectory (CBT) screw fixation exhibits distinct biomechanical characteristics compared with traditional pedicle screw (BPS) fixation. Specifically, CBT fixation reduces stress on the anterior column but concurrently increases stress on the posterior facet joints, an effect that is especially evident during rotational motion.\u003c/p\u003e"},{"header":"Abbreviations","content":"\u003cdiv class=\"DefinitionList\"\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eOLIF\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eOblique lumbar interbody fusion\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eASD\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eAdjacent segment degeneration\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eFE\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eFinite element\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eROM\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eRange of motion\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eLIF\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eLumbar interbody fusion\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eBPS\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003ebilateral pedicle screws\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eCBT\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003ecortical bone trajectory screws\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eIntact\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eintact osteoporosis model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L4 cage\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L4 cage model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L4\u0026thinsp;+\u0026thinsp;BPS\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L4 cage with bilateral pedicle screws model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L4\u0026thinsp;+\u0026thinsp;CBT\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L4 cage with bilateral cortical bone trajectory screws model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL4-L5 cage\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL4-L5 cage model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL4-L5\u0026thinsp;+\u0026thinsp;BPS\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL4-L5 cage with bilateral pedicle screws model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL4-L5\u0026thinsp;+\u0026thinsp;CBT\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL4-L5 cage with bilateral cortical bone trajectory screws model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L5 cage\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L5 cage model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L5\u0026thinsp;+\u0026thinsp;BPS\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L5 cage with bilateral pedicle screws model\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003cdiv class=\"DefinitionListEntry\"\u003e \u003cdiv class=\"Term\"\u003eL3-L5\u0026thinsp;+\u0026thinsp;CBT\u003c/div\u003e \u003cdiv class=\"Description\"\u003e \u003cp\u003eL3-L5 cage with bilateral cortical bone trajectory screws model.\u003c/p\u003e \u003c/div\u003e \u003c/div\u003e \u003c/div\u003e"},{"header":"Declarations","content":"\u003cp\u003e \u003cstrong\u003eEthics approval and consent to participate\u003c/strong\u003e \u003cp\u003eThe present study was approved by the Ethics Committee of the Third Hospital of Hebei Medical University. Informed consent obtained from each participant was written. All protocols are carried out in accordance with relevant guidelines and regulations.\u003c/p\u003e \u003c/p\u003e \u003cp\u003e \u003cstrong\u003eConsent for publication\u003c/strong\u003e \u003cp\u003eNot applicable.\u003c/p\u003e \u003c/p\u003e \u003cp\u003e \u003cstrong\u003eCompeting interests\u003c/strong\u003e \u003cp\u003eThe authors declare that they have no conflict of interest.\u003c/p\u003e \u003c/p\u003e\u003cp\u003e \u003ch2\u003eCompeting interests\u003c/h2\u003e \u003cp\u003eThe authors declare that they have no competing interests.\u003c/p\u003e \u003c/p\u003e\u003ch2\u003eFunding\u003c/h2\u003e\u003cp\u003eThis work was supported by the Natural Science Foundation of Hebei Province (H2022206429).\u003c/p\u003e\u003ch2\u003eAuthor Contribution\u003c/h2\u003e\u003cp\u003eConceptualization: Xianzhong Meng, Shuo Li, Lifei Wang, Hengrui ChangFormal Analysis: Xianzhong Meng, Shuo Li, Hengrui Chang, Kaibin Fan, Lifei Wang, Investigation: Jianhua Ren, Junkai Kou, Di ZhangMethodology: Shuo Li, Hengrui ChangProject Administration: Xianzhong MengWriting \u0026ndash; Original Draft: Shuo LiWriting \u0026ndash; Review \u0026amp; Editing: Shuo Li, Xianzhong Meng\u003c/p\u003e\u003ch2\u003eAcknowledgement\u003c/h2\u003e\u003cp\u003eWe would like to thank all the staff in Department of Spine Surgery, the Third Hospital of Hebei Medical University for their contribution on our research.\u003c/p\u003e\u003ch2\u003eData Availability\u003c/h2\u003e\u003cp\u003eThe datasets used and/or analyzed during the current study are available from the corresponding author on reasonable request.\u003c/p\u003e"},{"header":"References","content":"\u003col\u003e\u003cli\u003e\u003cspan\u003eXu C, Ji J, Zhang Y, et al. 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Global Spine J. 2022;12(5):851\u0026ndash;7.\u003c/span\u003e\u003c/li\u003e\u003c/ol\u003e"}],"fulltextSource":"","fullText":"","funders":[],"hasAdminPriorityOnWorkflow":false,"hasManuscriptDocX":true,"hasOptedInToPreprint":true,"hasPassedJournalQc":"","hasAnyPriority":false,"hideJournal":false,"highlight":"","institution":"","isAcceptedByJournal":false,"isAuthorSuppliedPdf":false,"isDeskRejected":"","isHiddenFromSearch":false,"isInQc":false,"isInWorkflow":false,"isPdf":false,"isPdfUpToDate":true,"isWithdrawnOrRetracted":false,"journal":{"display":true,"email":"[email protected]","identity":"bmc-musculoskeletal-disorders","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":false,"externalIdentity":"bmsd","sideBox":"Learn more about [BMC Musculoskeletal Disorders](http://bmcmusculoskeletdisord.biomedcentral.com/)","snPcode":"","submissionUrl":"https://author-welcome.nature.com/12891","title":"BMC Musculoskeletal Disorders","twitterHandle":"BMC_series","acdcEnabled":true,"dfaEnabled":true,"editorialSystem":"stoa","reportingPortfolio":"BMC Series","inReviewEnabled":true,"inReviewRevisionsEnabled":true},"keywords":"lumbar fusion, osteoporosis, finite element analysis, adjacent segment degeneration, cortical bone trajectory screws, pedicle screws, biomechanics","lastPublishedDoi":"10.21203/rs.3.rs-8272176/v1","lastPublishedDoiUrl":"https://doi.org/10.21203/rs.3.rs-8272176/v1","license":{"name":"CC BY 4.0","url":"https://creativecommons.org/licenses/by/4.0/"},"manuscriptAbstract":"\u003ch2\u003eBackground\u003c/h2\u003e \u003cp\u003eThis study investigated the biomechanical effects of single- and two-level oblique lateral interbody fusion (OLIF) with different internal fixation methods on adjacent segments in an osteoporotic lumbar spine model, focusing on range of motion (ROM), intervertebral disc stress, endplate stress, and facet joint contact forces.\u003c/p\u003e\u003ch2\u003eMethods\u003c/h2\u003e \u003cp\u003eA three-dimensional finite element (FE) model of the L1\u0026ndash;S1 spine was developed from CT scans of a healthy 30-year-old male and modified to simulate osteoporosis. The model included vertebral bodies, posterior elements, endplates, discs, and major ligaments. Material properties were assigned based on established literature, with modifications to simulate osteoporosis. Nine surgical scenarios were simulated: standalone cages, bilateral pedicle screws (BPS), and cortical bone trajectory (CBT) screws for both single- and two-level OLIF. A 400 N follower load and 10 Nm pure moments in six directions were applied to L1 to replicate physiological loading.\u003c/p\u003e\u003ch2\u003eResults\u003c/h2\u003e \u003cp\u003eAll fusion constructs increased ROM and mechanical stress at adjacent segments compared with the intact osteoporotic spine. Two-level fusion induced significantly greater biomechanical alterations than single-level fusion. In the L2\u0026ndash;L3 segment, the L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS model exhibited the highest flexion range of motion (ROM) of 6.70\u0026deg; and peak disc stress of 2.20 MPa, whereas the L3\u0026ndash;L4 cage\u0026thinsp;+\u0026thinsp;BPS configuration demonstrated a ROM of 6.07\u0026deg; and peak disc stress of 2.02 MPa. Similarly, at the L5\u0026ndash;S1 level, the L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS construct produced the greatest flexion ROM (14.69\u0026deg;) and disc stress (3.67 MPa), compared with the L4\u0026ndash;L5\u0026thinsp;+\u0026thinsp;BPS construct, which yielded a ROM of 13.14\u0026deg; and disc stress of 3.34 MPa. CBT fixation consistently produced lower disc and endplate stresses compared with BPS fixation; however, it resulted in substantially higher facet joint stresses, particularly during axial rotation. In the L3\u0026ndash;L5\u0026thinsp;+\u0026thinsp;CBT construct, L5\u0026ndash;S1 facet joint stress increased by 143% relative to the intact condition.\u003c/p\u003e\u003ch2\u003eConclusion\u003c/h2\u003e \u003cp\u003eTwo-level constructs impose greater cumulative loads than single-level ones. Compared with bilateral pedicle screws, cortical bone trajectory screws reduce anterior column loading but increase facet joint stress, especially during rotation.\u003c/p\u003e","manuscriptTitle":"Biomechanical Changes of Adjacent Segments Following Single- versus Two-level Oblique Lateral Interbody Fusion with Different Fixation Methods: A Finite Element study","msid":"","msnumber":"","nonDraftVersions":[{"code":1,"date":"2025-12-30 00:10:19","doi":"10.21203/rs.3.rs-8272176/v1","editorialEvents":[{"type":"communityComments","content":0},{"type":"decision","content":"Revision requested","date":"2026-01-27T08:39:34+00:00","index":"","fulltext":""},{"type":"editorInvitedReview","content":"","date":"2026-01-25T05:59:45+00:00","index":"hide","fulltext":""},{"type":"editorInvitedReview","content":"","date":"2026-01-18T21:43:17+00:00","index":"hide","fulltext":""},{"type":"editorInvitedReview","content":"","date":"2026-01-15T16:59:58+00:00","index":"hide","fulltext":""},{"type":"reviewerAgreed","content":"241227770255858252440745631737602079885","date":"2026-01-09T21:45:48+00:00","index":"hide","fulltext":""},{"type":"reviewerAgreed","content":"173528543688584373389564724074923182482","date":"2026-01-07T16:54:21+00:00","index":"hide","fulltext":""},{"type":"reviewerAgreed","content":"209505936453911466671243736496152563053","date":"2026-01-05T16:07:51+00:00","index":"hide","fulltext":""},{"type":"editorInvitedReview","content":"","date":"2025-12-26T16:03:49+00:00","index":"hide","fulltext":""},{"type":"reviewerAgreed","content":"148087191928523154039273096807385809107","date":"2025-12-18T10:33:26+00:00","index":"hide","fulltext":""},{"type":"reviewersInvited","content":"","date":"2025-12-18T06:40:40+00:00","index":"","fulltext":""},{"type":"editorInvited","content":"","date":"2025-12-04T15:22:33+00:00","index":"","fulltext":""},{"type":"editorAssigned","content":"","date":"2025-12-04T07:13:58+00:00","index":"","fulltext":""},{"type":"checksComplete","content":"","date":"2025-12-04T07:13:25+00:00","index":"","fulltext":""},{"type":"submitted","content":"BMC Musculoskeletal Disorders","date":"2025-12-03T16:05:17+00:00","index":"","fulltext":""}],"status":"published","journal":{"display":true,"email":"[email protected]","identity":"bmc-musculoskeletal-disorders","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":false,"externalIdentity":"bmsd","sideBox":"Learn more about [BMC Musculoskeletal Disorders](http://bmcmusculoskeletdisord.biomedcentral.com/)","snPcode":"","submissionUrl":"https://author-welcome.nature.com/12891","title":"BMC Musculoskeletal Disorders","twitterHandle":"BMC_series","acdcEnabled":true,"dfaEnabled":true,"editorialSystem":"stoa","reportingPortfolio":"BMC Series","inReviewEnabled":true,"inReviewRevisionsEnabled":true}}],"origin":"","ownerIdentity":"e31aadc7-0cf2-44d5-93f5-eb99d4176ed3","owner":[],"postedDate":"December 30th, 2025","published":true,"recentEditorialEvents":[],"rejectedJournal":[],"revision":"","amendment":"","status":"under-review","subjectAreas":[],"tags":[],"updatedAt":"2026-04-26T14:55:08+00:00","versionOfRecord":[],"versionCreatedAt":"2025-12-30 00:10:19","video":"","vorDoi":"","vorDoiUrl":"","workflowStages":[]},"version":"v1","identity":"rs-8272176","journalConfig":"researchsquare"},"__N_SSP":true},"page":"/article/[identity]/[[...version]]","query":{"redirect":"/article/rs-8272176","identity":"rs-8272176","version":["v1"]},"buildId":"XKTyCvWXoU3ODBz1xrDgd","isFallback":false,"isExperimentalCompile":false,"dynamicIds":[84888],"gssp":true,"scriptLoader":[]}