Axial rotation mechanics in a cadaveric lumbar ... - The Spine Journal

5 downloads 14803 Views 938KB Size Report
bDepartment of Neurosurgery and Orthopedics, College of Medicine, University of South Florida, 12901 Bruce B Downs Blvd, Tampa, FL 33620, USA.
The Spine Journal 14 (2014) 1272–1279

Basic Science

Axial rotation mechanics in a cadaveric lumbar spine model: a biomechanical analysis James J. Doulgeris, MSMEa,b,c,*, Sabrina A. Gonzalez-Blohm, MSBEa,b, Kamran Aghayev, MDa,b, Thomas M. Shea, BSa,b, William E. Lee, III, PhDd, Daniel P. Hess, PhDc, Frank D. Vrionis, MD, PhDa,b b

a NeuroOncology Program, H. Lee Moffitt Cancer Center & Research Institute, 13131 Magnolia Drive, Tampa, FL 33612, USA Department of Neurosurgery and Orthopedics, College of Medicine, University of South Florida, 12901 Bruce B Downs Blvd, Tampa, FL 33620, USA c Department of Mechanical Engineering, University of South Florida, 4202 E Fowler Ave, Tampa, FL 33620, USA d Department of Chemical & Biomedical Engineering, University of South Florida, 4202 E Fowler Ave, Tampa, FL 33620, USA Received 30 May 2013; revised 18 October 2013; accepted 21 November 2013

Abstract

BACKGROUND CONTEXT: Postoperative patient motions are difficult to directly control. Very slow quasistatic motions are intuitively believed to be safer for patients, compared with fast dynamic motions, because the torque on the spine is reduced. Therefore, the outcomes of varying axial rotation (AR) angular loading rate during in vitro testing could expand the understanding of the dynamic behavior and spine response. PURPOSE: To observe the effects of the loading rate in AR mechanics of lumbar cadaveric spines via in vitro biomechanical testing. STUDY DESIGN: An in vitro biomechanical study in lumbar cadaveric spines. METHODS: Fifteen lumbar cadaveric segments (L1–S1) were tested with varying loading frequencies of AR. Five different frequencies were normalized with the base line frequency (0.125 Hz n515) in this analysis: 0.05 Hz (n56), 0.166 Hz (n56), 0.2 Hz (n510), 0.25 Hz (n510), and 0.4 Hz (n58). RESULTS: The lowest frequency (0.05 Hz) revealed significant differences (p!.05) for all parameters (torque, passive angular velocity, axial velocity [AV], axial reaction force [RF], and energy loss [EL]) with respect to all other frequencies. Significant differences (p!.05) were observed in the following: torque (0.4 Hz with respect to 0.2 Hz and 0.25 Hz), passive sagittal angular velocity (SAV) (0.4 Hz with respect to all other frequencies; 0.166 Hz with respect to 0.25 Hz), axial linear velocity (0.4 Hz with respect to all other frequencies), and RF (0.4 Hz with respect to 0.2 Hz and 0.25 Hz). Strong correlations (R2O0.75, p!.05) were observed between RF with intradiscal pressure (IDP) and AR angular displacement with IDP. Intradiscal pressure (p!.05) was significantly larger in 0.2 Hz in comparison with 0.125 Hz. CONCLUSIONS: Evidences suggest that measurements at very small frequencies (0.05 Hz) of torque, SAV, AV, RF, and EL are significantly reduced when compared with higher frequencies (0.166 Hz, 0.2 Hz, 0.25 Hz, and 0.4 Hz). Higher frequencies increase torque, RF, passive SAV, and AV. Higher frequencies induce a greater IDP in comparison with lower frequencies. Ó 2014 Elsevier Inc. All rights reserved.

Keywords:

Range of motion; Energy loss; Loading rate; Intradiscal pressure; Passive motion; Axial displacement

FDA device/drug status: Not applicable. Author disclosures: JJD: Nothing to disclose. SAG-B: Nothing to disclose. KA: Nothing to disclose. TMS: Nothing to disclose. WEL: Nothing to disclose. DPH: Nothing to disclose. FDV: Grants: Globus Spine (E, Paid directly to institution), Synthes Spine (E, Paid directly to institution), Orthofix (C, Paid directly to institution), Backbone Medical (C, Paid directly to institution). 1529-9430/$ - see front matter Ó 2014 Elsevier Inc. All rights reserved. http://dx.doi.org/10.1016/j.spinee.2013.11.037

The disclosure key can be found on the Table of Contents and at www. TheSpineJournalOnline.com. * Corresponding author. NeuroOncology Program, H. Lee Moffitt Research Center, 13131 Magnolia Drive, Tampa, FL 33612, USA. Tel.: (813) 745-6084; fax: (813) 745-3917. E-mail address: [email protected] (J.J. Doulgeris)

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

Introduction

Methods

Back pain disorders are among the most common causes for work disability, and ‘‘twisting while lifting or carrying’’ has been identified as one of the leading causes for back sprain and strain injuries [1]. Axial rotation (AR) is one of the basic motions associated with average daily living movements of the spine, such as walking [2], and is more strenuously applied in sports, such as golf and tennis [3]. A literature review of golf swing biomechanics conducted by Gluck et al. [3] emphasizes how motions performed during this sport are commonly underestimated and their complexity can easily trigger spinal injury. In fact, Gluck et al. [3] concluded the human body may not be designed to endure forces like the ones generated by swinging. Injuries and low back pain have motivated researchers to investigate AR mechanics [4–6]. Very slow quasistatic motions are intuitively believed to be safer for patients, compared with fast dynamic motions, because the torque on the spine is reduced. However, patients may be more inclined to move faster after a surgery because pain may be mitigated. Faster loading rates increase stresses on the spine that may risk injury, compromise surgical success, and possibly accelerate adjacentsegment degeneration. Investigation of the loading rate may provide insight to postoperative care. Axial rotation investigations focus on kinematics and dynamics that assist in the explanation of some possible problems. Biomechanical models, based on in vivo trunk kinematics, suggest that the trunk’s capacity can be easily exceeded during daily torsional exertion even at low velocities [4]. Additionally, dynamic motions can compromise the ability to apply twisting motions by 50% [7]. Furthermore, Farfan et al. [5] suggested a possible relationship between disc injury, torsional stresses, and disc degeneration when experimentally-induced disc ruptures triggered similar effects to those due to degeneration. Also, Farfan et al. [5] suggested a possible relationship between annulus damage during twisting injury and ‘‘bulging’’ disc condition. The in vivo effects of AR motion on spine biomechanics have previously been described by two parameters: intervertebral rotation and intradiscal pressure (IDP) [8]. The inclusion of these two variables in the characterization of axial motion for in vitro biomechanical testing is well accepted. The rate of trunk axial moment application has been shown to influence the relative activity of multiple trunk muscles during in vivo investigations [4]. Other studies have indicated or concluded that the loading rate can increase the risk of failure in flexion and extension motions [9,10]. Therefore, the outcomes of varying AR angular loading frequencies during in vitro testing could expand the understanding of the dynamic behavior and spine response. The aforementioned reasons motivated our research and defined the aim of the study: to observe the effects of the loading rate in AR mechanics of lumbar cadaveric spines via in vitro biomechanical testing.

Specimen preparation

1273

Fifteen fresh cadaveric lumbar spines containing 12 males and 3 females (average age of 54 years, standard deviation of 6) were used in this investigation. Specimens were dissected into L1–S1 segments and proper care was taken to preserve the intervertebral discs, synovial capsules, and ligaments. Specimens were thawed in a refrigerator (463 C) overnight before dissection and testing. Gauze sponges (400 400 ) were wrapped around all the exposed tissue and moistened with 0.9% NaCl solution when specimens were not being tested. Specimens were kept out of a frozen environment for a maximum of 48 hours for dissection and testing sequence. Self-tapping screws (200 ) were installed into the superior (L1) and inferior (S1) vertebral bodies, then potted into a custom frame via a polyester resin (Bondo; Bondo Corp, Atlanta, GA, USA). Leveling and a customized potting frame alignment tools were used to ensure that vertebral bodies were centered, frames were parallel, and the frames were not offset.

Testing machine Specimens were loaded into a custom servo hydraulic testing apparatus (MTS 858 MiniBionix; modified by Instron, Norwood, MA, USA) to apply controlled forces and motions. The machine was customized to a 4 df system that allows (1 and 2) bending motion on both superior and inferior frames, (3) AR, and (4) displacement on the superior frame. The inferior frame remained fixed from both AR and displacement. The customized testing apparatus has been used and accepted in previous publications [11]. Global (L1–S1) angular (60.1 ) and linear (60.1 mm) axial displacements were measured by the hydraulic actuator’s internal sensors (MTS 242), whereas torque (60.1 Nm) and axial reaction force (RF) on the specimen (61 N) were measured by a load cell (Dynacell, MTS). Angular and axial displacements were also tracked optoelectronically by sensors located at the superior and inferior frames through an Optotrak Certus system (Optotrak 3020; Northern Digital, Inc., Waterloo, Canada). Intradiscal pressure (60.01 MPa) was measured in six of the specimens at L2–L3 using a 060S pressure transducer (Precision Measurement Company, Ann Arbor, MI, USA) [2,12]. The pressure transducers were inserted into the center of the nucleus pulposus and were not removed between tests. The center of the intervertebral disc was targeted by caliper measurements of the axial, coronal, and sagittal planes. A cannulated needle was marked and inserted laterally so the pressure transducer could be inserted into the needle until it reached the middle of the nucleus. After inserting the pressure transducer, the needle was removed. Transducer location and disc conditions were verified and examined ex post facto. L2–L3 was chosen because it is superior to

1274

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

a common surgical fusion level (L3–L4) and will be used for comparison in future investigations. Pressure transducer signals were conditioned by a signal conditioner (System 2100; Vishay Micro Measurements, Wendell, NC, USA) and IDP was recorded by the Optotrak data acquision system. All data were acquired at a rate of 10 Hz. Biomechanical protocol Specimens were placed under a 50.0 N axial preload [13] before all AR tests. The specimens underwent axial torque controlled conditioning cycles (6 cycles of 65 Nm at 0.125 Hz) so that the angular limits could be determined. The angular limits from the last three torque controlled cycles were averaged and used as the amplitude in a position-controlled sinusoidal function, whereas frequency was varied. Position control, unlike torque control, does not require a feedback loop to drive the motor, which makes the conditions more repeatable when loading frequency is varied. Five different frequencies were normalized with the base line frequency (0.125 Hz, n515) in this analysis: 0.050 Hz (n56), 0.166 Hz (n56), 0.200 Hz (n510), 0.250 Hz (n510), and 0.400 Hz (n58). Tests were repeated on three random specimens to determine the relative error in the measurements and confirm that no plastic deformation occurred. A summary of the tests conducted is presented in Table 1. Analysis Axial rotation mechanics were expressed by measuring the following variables at L1: torque, AR range of motion (ROM), axial velocity (AV), reaction force, axial displacement (AD), and IDP. Sagittal (flexion/extension [FE]) angular velocity (SAV) was measured with respect to L1 and S1 angular displacements. Intradiscal pressure was performed on Table 1 Schematic representation of conditions and the number of sets performed in each specimen Frequencies tested 0.125 Hz Specimen (baseline)

0.050 0.166 0.200 0.250 0.400 Intradiscal Hz Hz Hz Hz Hz pressure

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 Total515

X X X X X

X X X X X X X X X X X XXX XX XXX X n515

X

X X X X X X

X n56

n56

X X X X X X

Fig. 1. Equations used to estimate energy loss, passive sagittal angular velocity, and passive axial velocity. (Top) Definition of energy loss of a viscously damped system per cycle. (Middle) Derivation of the trapezoidal integration equation for a hysteresis loop. (Bottom) Average velocity and slope equation. Where DE is energy loss, Fd is the force damped, x is the displacement (angular or axial), n is the number of frames, T is torque, q is the angular displacement, V is velocity, and t is time.

0.125 Hz, 0.166 Hz, and 0.200 Hz, whereas all other measurements included the full range of frequencies (Table 1). Energy loss (EL) of a viscously damped system was described by Inman [14] as the area enclosed by the hysteresis loop (Fig. 1, Top). The EL was estimated by trapezoidal numerical integration (Fig. 1, Middle) of the area enclosed by the hysteresis loop (Fig. 2). Sagittal angular velocity and AV were estimated with basic differentials (Fig. 1, Bottom) using linear regression of the loading motion (Fig. 3). AD and FE ROM measurements were shifted to correct phase lag with the other measurements. All measurements were normalized with respect to 0.125 Hz to remove interspecimen effects. Intradiscal pressure was excluded from normalization because frequencies were compared with respect to the same specimens. Normal distribution of all parameters was verified with a Shapiro-Wilk test. A repeated measure analysis of variance and a multicomparison posthoc test, with a Bonferroni correction, were used to determine the significant differences among the frequencies. Linear regression was used to determine correlations between measurements and IDP during loading of the specimen. Only loading was observed for IDP analysis to reduce the error from plastic deformation

X X X X X

X X X X X X XXX XXX XXX XX XX XX XXX XXX XXX X X n510 n510 n59

X X X X X X

n56

Note: X, specimen tested under the corresponding variable for X sets of six cycles. For example, XX represents 12 cycles of axial rotation.

Fig. 2. Representation of the motion that the specimen takes while being loaded under the baseline frequency (hysteresis loop).

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

1275

Table 2 Schematic representation of statistical differences for all variables with respect to the different frequencies tested Comparisons

Torque

SAV

AV

RF

EL

D D D D

D D D D

D D D D

D D D D

D D D D

D D

D

D

D

D

D

D

D

D

D

Fig. 3. Representation of one cycle of the global flexion/extension, axial displacement, axial rotation displacement, and torque at 0.125 Hz. Linear regressions (black lines) represent the slope from which the passive sagittal angular velocity and axial velocity averages were estimated.

0.050 Hz with 0.166 Hz 0.200 Hz 0.250 Hz 0.400 Hz 0.166 Hz with 0.200 Hz 0.250 Hz 0.400 Hz 0.200 Hz with 0.250 Hz 0.400 Hz 0.250 Hz with 0.400 Hz

when directions were changed. A significance level of 0.05 was used for all statistical analysis.

SAV, sagittal angular velocity; AV, axial velocity; RF, reaction force; EL, energy loss. D, significantly different (p!.05).

Results Statistical results for torque, SAV, AV, RF, IDP, and EL are summarized in Table 2 and normalized data are displayed in Figs. 4 and 5. The lowest frequency (0.05 Hz) revealed significant differences (p!.05) for all parameters (torque, SAV, AV, RF, and EL) with respect to all other frequencies. Additionally, there was also enough evidence to state the following significant differences (p!.05): torque (0.4 Hz with respect to 0.2 Hz and 0.25 Hz), SAV (0.4 Hz with respect to all other frequencies; 0.166 Hz with respect to 0.25 Hz), AV (0.4 Hz with respect to all other frequencies), and RF (0.4 Hz with respect to 0.2 Hz and 0.25 Hz). Strong correlations (R2O0.75, p!.05) were observed between RF-IDP and AR ROM-IDP (Fig. 6). Medium

correlations (0.50!R2!0.75, pO.05) were observed between AD-IDP and FE ROM-IDP (Fig. 6). Range of motion, AV, and torque had a relative error of less than 2%, whereas EL and SAV had a relative error of less than 4% (n53). The sizes of the relative errors and the lack of drift lead to the assumption that minimal plastic deformation occurred throughout the tests. Discussion The facets, ligaments, and discs are the main contributors in spine structure and mechanics [15]. Nevertheless, the facets and ligaments provide most of the rigid structural strength in AR [5] (supporting the lack of drift observed in the repeated tests), whereas the intervertebral discs provide a

Fig. 4. Normalized results with respect to baseline frequency, 0.125 Hz, for reaction force, torque, energy loss, intradiscal pressure, passive sagittal angular velocities, and passive axial velocities. Error bars indicate the standard deviation of the measurements. Symbols (a, b, g, d, ε, U) represent significant differences (p!.05) with respect to 0.050 Hz (a), 0.166 Hz (b), 0.200 Hz (g), 0.250 Hz (d), 0.400 Hz (ε), all other frequencies (U).

1276

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

Fig. 5. Intradiscal pressure results during axial rotation. The asterisk indicates significant difference (p!.05). Error bars indicate the standard error of the measurements.

smooth transition through the neutral zone [16]. Hypothetically, if the discs were main contributors to structural strength, then it would be expected that the creep or drift would be apparent, considering the amount of cycles. The discs contribute the most, in AR, when the direction is changing from one to another (neutral zone) and the least when the facets are contacting and ligaments are taut (elastic zone). Coupled motions Panjabi [16] described the mechanics of the spine by introducing the concepts of the neutral and elastic zones. However, AR neutral and elastic zones are different than

FE and lateral bending because the other motions have larger neutral zones and do not significantly shear the disc. The AR neutral zone occurs in the ‘‘backlash area’’ of the hysteresis loop and has no facet interaction and the elastic zone is located in all other areas where facets are interacting (Fig. 2). The AR elastic zone is larger than the neutral zone (Fig. 2) and was further described by two subsections: transitional and structural. The transitional section is characterized when the facets are sliding and inducing passive motions/torques in other directions, whereas the structural section is described when the posterior ligaments restrict passive motions/forces and facet sliding. The coupling motions in the transitional section are controlled by the anatomical mechanics of the spine and are directly related to the angulation of the facet joints with respect to the sagittal plane. The coupling effects may result in motion or residual torque during AR and are the reason the SAV and AD occur during dynamic rotation. The passive coupled motion results support the in vivo sagittal coupled moments observed during torsional motion [4,17]. The coupled motions in each of the specimens created passive flexion during loading and passive extension during unloading (Fig. 3). The passive FE rotation responses were attributed to the rostral displacement in posterior column caused from facet surface interaction during AR. Passive FE rotation is an attempt to strain under the path of least resistance and suggests that forcing flexion should reduce AR stiffness. The previous explanation is validated by an

Fig. 6. Representation of linear regressions performed to establish correlations between intradiscal pressure, passive sagittal angular rotation, axial displacement, and axial rotation.

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

in vivo investigation that observed that AR ROM increases under forced flexion [18,19]. Alternatively, an in vitro investigation did not have enough evidence to warrant a difference in AR ROM between unloaded and loaded flexions when testing functional spinal units, but did observe a significant decrease in AR ROM under forced extension [20]. The extent of forced loading may only significantly affect global measurements (full lumbar) and have minimal impact on functional segments. Intradiscal pressure The increases in IDP are mainly because of shear on the disc, but also have some components from the passive flexion and axial compression during AR. The IDP profiles slightly lagged (less than 0.2 seconds) behind the driving motion with a maximum peak near full rotation and a minimal peak at the neutral position (equilibrium). The results confirm that IDP is highly correlated with ROM (Fig. 6) and validates evidence from other studies [21]. The magnitudes of the IDP measured at L2–L3 (Fig. 6) are in the range of in vitro and in vivo studies measured at L3–L4 and L4–L5 [22–25]. The weaker correlation in passive FE rotation indicates that the AR rotation (disc shear) may have a higher influence on IDP (Fig. 6). Significant IDP differences were observed in certain frequencies that were not observed with respect to RF (Table 2), which suggests that local IDP is more sensitive than the relative (global) RF measurements. The variation of IDP and RF triggered by changes in angular frequency supports that the disc is sensitive to the loading frequency [26]. Lastly, the strong correlation of RF and IDP (Fig. 6) indicate that maximum pressure may occur at higher AR velocities. The findings corroborate that faster movements may trigger higher stresses than similar slower movements [9,10]. Axial torque Wilke et al. [27] postulated how to measure stiffness in the neutral and elastic zones to describe the innate phenomena of nonlinear stiffness of the spine that is observed in the hysteresis loop. The loading rates have no effect on stiffness but will affect the torque profile. Angular velocity, damping coefficient, and acceleration alter the torque profile in a damped system [14]. Stiffness, at a given displacement, is best calculated at zero acceleration (quasistatic) or approximated at very low frequencies (#0.05 Hz) because the damping and acceleration forces are minimized. Higher loading rates, regardless of the extent of the ROM, increase and concentrate the stresses on the spine. Stress concentrations are dependent on the extent of motion and anatomy, but cause failure when torque is increased. Investigations have concluded that torque or risk of injury will increase with respect to loading rate in FE [9,10], which supports the findings of increased AR torque with respect to loading rate. The moment of area that affects an in vitro spine is smaller in comparison with an in vivo body because of

1277

the lack of upper torso and extremity incorporation. Furthermore, the in vivo moment of area, with respect to AR, is dependent on extremity position (complete arm adduction or medial arm abduction) and midsection weight distribution. In vivo AR loading rates may have higher peak torques when compared with in vitro results because larger moment of area will create a residual momentum when rotated. Also, the in vivo response will depend on the natural frequency of the spine. A lumbar spine in vitro natural frequency investigation has been performed [28], but the variability of in vivo moment of area makes an in vivo extrapolation very difficult. Therefore, the lower moment of areas used in this study limit the extrapolation and underestimate the peak torque during increased loading rates. Energy loss Energy can be dissipated in the form of heat, permanent deformation, and/or damping and has been mentioned in other publications [27]. Energy loss in spine biomechanics is a concept that is not normally included in investigations, but incorporation adds depth to the analysis. For example, EL at low frequencies (0.05 Hz) estimates the overall plasticity in the specimen and EL measurements at higher frequencies will estimate damping, plasticity, and other losses. The overall plasticity (EL estimation at very low frequencies) is best defined by an evolved neutral zone explanation as described by Panjabi [16]. The neutral zone is estimated by measuring the residual displacement after loading [16,27]. Neutral zone measurements describe the magnitude of residual displacement, where overall relative plasticity describes the entire residual area. The residual area can quantify the elasticity of a specimen and will have an impact in fusion analyses. Other studies have shown that AR EL can be used for biomechanical comparisons [29]. Plastic deformation is more likely to occur in the disc because it has a lower modulus of elasticity and yield point compared with the vertebrae. Therefore, in vitro spine EL is related to the amount of disc contribution during rotations. The sensitivity of EL with respect to frequency indicates the amount of disc contribution with respect to motion because the disc is the main source of damping in the in vitro spine. The results show that the amount of EL was significantly smaller and more reproducible in low frequencies (0.05 Hz), but no other differences were observed among the other frequencies (0.166–0.4 Hz). The lack of difference, between the frequencies 0.166 Hz and 0.4 Hz, indicates that the damping EL is minimal and AR is more dependent on the rigid structures. The results of the investigation suggest that the loading rate should be low as possible to reduce the measurement error for future implant investigations. Clinical implications Retraining muscle motions and haptic feedback interpretation after a fusion may prevent concentrated stresses on the spine after surgery, but the continuation of presurgical

1278

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

motion memories may have the adverse effect. Muscle contributions in AR have been investigated by others [4,30– 32], but haptic feedback is inherent and subconscious. Haptic feedback and antagonizer muscles ensure motions do not overexert the spine, but safety is compromised with altered kinematics at higher frequencies. Furthermore, studies have clinically investigated or reviewed the relationship between adjacent-segment degeneration and fusion, but no conclusive evidence guarantees adjacent-segment degeneration after fusion [33–35]. Fusion procedures reduce ROM and may generate faster loading rates above and below a segmented fusion in compensation for the altered dynamics; however, this hypothesis must be investigated. Haptic feedback governs the speed of spine motions when pain is involved. Typically, pain causes a patient to move slower to prevent additional harm to the spine. The general goal of spine surgery is to reduce pain as much and as quickly as possible, but this may cause patients to move faster after surgery. Lumbar surgical success could rely on the prevention of excessive torques that the spine endures during higher loading rates. Thus, patients may benefit from physical therapy before and after fusion to prevent increased loading rates, but this requires insurance approval and time to take effect. Therefore, lumbar surgical patients, especially if in an outpatient capacity, may be advised to inflate the diaphragm before all twisting motions to help prevent adverse loading rates immediately after surgery. Extrapolation of the results of this study, to the posttraumatic spine, is difficult because the anatomy of the specimens remained intact. The segmental AR stiffness, of a full intact model, is similar, but a segment’s stiffness can be altered after trauma in comparison with the surrounding segments. Large shifts in segmental stiffness in a multielement system can unbalance the load distribution and become sensitive to loading rates. In general, a posttraumatic spine should not undergo extensive loading or fast loading rates because it could risk more trauma and compromise alignment, but this depends on the extent of the trauma. Further investigations are required to understand how an injury can affect the kinematics and dynamics of a spine loaded under different frequencies. Comparison of fusion constructs is a topic of great interest for research investigations and used by spine surgeons to review implant performance before surgical installment. In vitro investigations, in general, measure ROM and stiffness of functional spinal units and use them to compare constructs. The inclusion of the neutral zone measurement has gained popularity among investigations, but small differences between conditions can make it difficult to find significant differences within small sample sizes. Conversely, EL is more sensitive to alterations in the spine. Energy loss is an estimation of the disc’s plastic (permanent) deformation during motions. Therefore, EL has a strong impact on fusion analyses and can help determine how a device will perform in clinical scenarios [29].

Limitations Several limitations can be mentioned in this investigation. The use of cadaveric spines may provide a source of error in the results and interpretation. The torque was delivered to the top of the spine (L1), but the muscles deliver torques throughout the spine. Intradiscal pressures were measured with a pressure transducer, which is highly sensitive to disc dehydration and motion. All significant plastic deformation was assumed to happen during the precondition torque controlled cycle. The sample size was limited because of donor unavailability. Conclusion The investigation demonstrated the sensitivity of RF, IDP, and AR mechanics with respect to frequency. Evidences suggest that measurements at very small frequencies (0.05 Hz) of torque, SAV, AV, RF, and EL are significantly reduced when compared with higher frequencies (0.166 Hz, 0.2 Hz, 0.25 Hz, and 0.4 Hz). Higher frequencies increase torque, RF, SAV, and AV with respect to lower frequencies. Higher frequencies (0.2 Hz) have a larger IDP than lower frequencies (0.125 Hz). We recommend that EL be measured at low velocities to reduce the error of the measurements. Acknowledgments The authors would like to thank Beth DeMarse for her assistance in physical therapy consultation. No direct funding was used to support this investigation, nor was there any conflict of interest. References [1] Labor USDo. Back disorders and injuries. Available at: http://www. osha.gov/dts/osta/otm/otm_vii/otm_vii_1.html#3. Accessed December 12, 2012. [2] Cannella M, Arthur A, Allen S, et al. The role of the nucleus pulposus in neutral zone human lumbar intervertebral disc mechanics. J Biomech 2008;41:2104–11. [3] Gluck GS, Bendo JA, Spivak JM. The lumbar spine and low back pain in golf: a literature review of swing biomechanics and injury prevention. Spine J 2008;8:778–88. [4] Marras WS, Granata KP. A biomechanical assessment and model of axial twisting in the thoracolumbar spine. Spine 1995;20:1440–51. [5] Farfan HF, Cossette JW, Robertson GH, et al. The effects of torsion on the lumbar intervertebral joints: the role of torsion in the production of disc degeneration. J Bone Joint Surg Am 1970;52:468–97. [6] Burnett A, O’Sullivan P, Ankarberg L, et al. Lower lumbar spine axial rotation is reduced in end-range sagittal postures when compared to a neutral spine posture. Man Ther 2008;13:300–6. [7] Cho KS, Kang SG, Yoo DS, et al. Risk factors and surgical treatment for symptomatic adjacent segment degeneration after lumbar spine fusion. J Korean Neurosurg Soc 2009;46:425–30. [8] Dreischarf M, Rohlmann A, Bergmann G, Zander T. Optimised loads for the simulation of axial rotation in the lumbar spine. J Biomech 2011;44:2323–7. [9] Wang JL, Parnianpour M, Shirazi-Adl A, Engin AE. Viscoelastic finite-element analysis of a lumbar motion segment in combined

J.J. Doulgeris et al. / The Spine Journal 14 (2014) 1272–1279

[10]

[11]

[12]

[13]

[14] [15]

[16] [17]

[18]

[19] [20]

[21]

[22]

compression and sagittal flexion. Effect of loading rate. Spine 2000; 25:310–8. Adams MA, Dolan P. Recent advances in lumbar spinal mechanics and their clinical significance. Clin Biomech (Bristol, Avon) 1995;10: 3–19. Setzer M, Eleraky M, Johnson WM, et al. Biomechanical comparison of anterior cervical spine instrumentation techniques with and without supplemental posterior fusion after different corpectomy and discectomy combinations: laboratory investigation. J Neurosurg Spine 2012;16:579–84. Tzermiadianos MN, Renner SM, Phillips FM, et al. Altered disc pressure profile after an osteoporotic vertebral fracture is a risk factor for adjacent vertebral body fracture. Eur Spine J 2008;17:1522–30. Slucky AV, Brodke DS, Bachus KN, et al. Less invasive posterior fixation method following transforaminal lumbar interbody fusion: a biomechanical analysis. Spine J 2006;6:78–85. Inman DJ. Engineering vibration. Upper Saddle River, NJ: Prentice Hall, 2008. Panjabi MM. The stabilizing system of the spine. Part I. Function, dysfunction, adaptation, and enhancement. J Spinal Disord 1992;5: 383–9; discussion 397. Panjabi MM. The stabilizing system of the spine. Part II. Neutral zone and instability hypothesis. J Spinal Disord 1992;5:390–6; discussion 397. Ng JK, Parnianpour M, Richardson CA, Kippers V. Functional roles of abdominal and back muscles during isometric axial rotation of the trunk. J Orthop Res 2001;19:463–71. Drake JDM, Callaghan JP. Do flexion/extension postures affect the in vivo passive lumbar spine response to applied axial twist moments? Clin Biomech 2008;23:510–9. Pearcy MJ. Twisting mobility of the human back in flexed postures. Spine 1993;18:114–9. Haberl H, Cripton PA, Orr TE, et al. Kinematic response of lumbar functional spinal units to axial torsion with and without superimposed compression and flexion/extension. Eur Spine J 2004;13:560–6. Heuer F, Schmidt H, Claes L, Wilke HJ. Stepwise reduction of functional spinal structures increase vertebral translation and intradiscal pressure. J Biomech 2007;40:795–803. Wilke HJ, Wolf S, Claes LE, et al. Influence of varying muscle forces on lumbar intradiscal pressure: an in vitro study. J Biomech 1996;29: 549–55.

1279

[23] Wilke HJ, Neef P, Caimi M, et al. New in vivo measurements of pressures in the intervertebral disc in daily life. Spine 1999;24:755–62. [24] Schmoelz W, Huber JF, Nydegger T, et al. Influence of a dynamic stabilisation system on load bearing of a bridged disc: an in vitro study of intradiscal pressure. Eur Spine J 2006;15:1276–85. [25] Sato K, Kikuchi S, Yonezawa T. In vivo intradiscal pressure measurement in healthy individuals and in patients with ongoing back problems. Spine 1999;24:2468–74. [26] Iatridis JC, Setton LA, Weidenbaum M, Mow VC. The viscoelastic behavior of the non-degenerate human lumbar nucleus pulposus in shear. J Biomech 1997;30:1005–13. [27] Wilke HJ, Wenger K, Claes L. Testing criteria for spinal implants: recommendations for the standardization of in vitro stability testing of spinal implants. Eur Spine J 1998;7:148–54. [28] Crisco JJ, Fujita L, Spenciner DB. The dynamic flexion/extension properties of the lumbar spine in vitro using a novel pendulum system. J Biomech 2007;40:2767–73. [29] Doulgeris JJ, Aghayev K, Gonzalez-Blohm SA, et al. Comparative analysis of posterior fusion constructs as treatments for middle and posterior column injuries: an in vitro biomechanical investigation. Clin Biomech (Bristol, Avon) 2013;28:483–9. [30] Ng JK, Richardson CA, Parnianpour M, Kippers V. EMG activity of trunk muscles and torque output during isometric axial rotation exertion: a comparison between back pain patients and matched controls. J Orthop Res 2002;20:112–21. [31] Kumar S, Narayan Y, Zedka M. An electromyographic study of unresisted trunk rotation with normal velocity among healthy subjects. Spine 1996;21:1500–12. [32] Mirka G, Kelaher D, Baker A, et al. Selective activation of the external oblique musculature during axial torque production. Clin Biomech (Bristol, Avon) 1997;12:172–80. [33] Cheh G, Bridwell KH, Lenke LG, et al. Adjacent segment disease following lumbar/thoracolumbar fusion with pedicle screw instrumentation: a minimum 5-year follow-up. Spine 2007;32:2253–7. [34] Hilibrand AS, Robbins M. Adjacent segment degeneration and adjacent segment disease: the consequences of spinal fusion? Spine J 2004;4(6 Suppl):190S–4S. [35] Throckmorton TW, Hilibrand AS, Mencio GA, et al. The impact of adjacent level disc degeneration on health status outcomes following lumbar fusion. Spine 2003;28:2546–50.