Functional spine unit development

Transcription

1 Functional spine unit development Literature review Ed Fournier Michael Wonnacott Biokinetics and Associates Ltd Don Reid Drive Ottawa, Ontario K1H 1E1 PWGSC Contract Number: W /001/QCL (TA84) CSA: Robert Durocher The scientific or technical validity of this Contract Report is entirely the responsibility of the Contractor and the contents do not necessarily have the approval or endorsement of the Department of National Defence of Canada. This document was reviewed for Controlled Goods by DRDC using the Guide to Canada s Export Controls. Contract Report DRDC-RDDC-2016-C161 May 2016

2 Her Majesty the Queen in Right of Canada, as represented by the Minister of National Defence, 2016 Sa Majesté la Reine (en droit du Canada), telle que représentée par le ministre de la Défense nationale, 2016

3 Significance for Defence and Security Vehicles subjected to IED detonation exposes the occupants to significant high rate loading in the vertical direction. Seating systems designed to mitigate the magnitude of the forces transmitted to the occupants have been implemented. The evaluation of these seats is carried out in full-scale blast tests using automotive crash test dummies. The lumbar spine of these dummies were not designed for loading in the vertical direction and consequently do not properly assess the loads transmitted to the occupant. A new lumbar spine element is being developed such that loading and threat of injury can be better evaluated. i

4 Abstract... The prevalence of lumbar spine injuries from underbody blast (UBB) has been noted to result in debilitating long terms effects on mounted servicemen and women. The energy absorbing vehicle seats that have been implemented do provide some protection against UBB loading but the level of protection cannot be properly assessed using current test methods. The Hybrid III anthropomorphic test device (ATD) used in experimental blast trials along with the Dynamic Response Index (DRI) injury criteria for assessing lumbar spine injuries are believed to be inadequate for assessing the various modes of dynamic loading and related injury risks. A literature review was conducted to identify the parameters that may affect a surrogate s ability to assess injury. These include the effect of seating posture and increased inertial loads from additional body borne mass on the loads applied to the lumbar spine during an UBB incident. Biofidelic characterization of the functional spine unit was sought for the development of an improved functional spine unit for assessing spine injury. Of principal concern is an ATD s ability to be correctly oriented during initial placement and have the required sensing capabilities to monitor for combined loading conditions commensurate with the selected injury assessment reference values. ii

5 Table of contents Significance for Defence and Security... i Abstract ii Table of contents... iii List of figures... iv List of tables... v 1 Introduction Problem Definition Effects of Seating Posture Upper Body Mass Effects On Lumbar Injury Spine Response to Loading Response to Compression Loading Response to Shear Loading Response to Flexion-Shear Loading Range of Motion Existing Injury Criteria Eiband Injury Tolerance Curve for Vertical Acceleration Dynamic Response Index Multi-axial Dynamic Response Criteria Lumbar Load Criterion Spine Injury Criterion Anthropomorphic Test Devices Frontal Impact Dummies Hybrid III THOR-M Rear Impact Dummies RID3D A.1.1 BioRID II Side Impact Dummies Other Dummies ADAM Dummy WIAMan Surrogate Spine Development Approach Summary and Conclusion References/Bibliography References Bibliography List of symbols/abbreviations/acronyms/initialisms iii

6 List of figures Figure 1: Possible loading mechanism that may result in fracture (figure from [3]) Figure 2: Mean translational displacements of lumbar spine segments under preload conditions Figure 3: Mean rotational displacements of lumbar spine segments under preload conditions Figure 4: Compressive force deflection curve for; (a) normal, and; (b) degenerated intervertebral disk loaded in compression (figure from [14]) Figure 5: Probability curves for spinal injury due to compressive loading (figure from [15]) Figure 6: Compressive stiffness of intervertebral disks by lumbar level and loading rate [16] Figure 7: Typical shear results for a L2-L3 motion segment (figure from [17]) Figure 8: Mean flexion, extension, lateral bending and torsion response of motion segments tested by Schultz et al [18]. Dashed lines are tests with the posterior elements removed Figure 9: Mean compression, lateral, anterior and posterior shear response of motion segments tested. Dashed lines are tests with the posterior elements removed (figure from [19]) Figure 10: Flexion and versus bending moment at failure of lumbar motion segments (figure from [20]) Figure 11: Typical bending moment versus time response for severe load with a 15 ms and a 5 ms pulse rise time (figure from [23]) Figure 12: Moment deflection curve up to the elastic limit of the FSU (figure from [25]) Figure 13: Eiband injury tolerance curve for vertical acceleration (figure from [32]) Figure 14: DRI lumped mass model of the spine (figure from [33]) Figure 15: DRI versus spinal injury rate (figure from [32]) Figure 16: Critical point for computing BDRC (figure from [36]) Figure 17: Correlation between DRI and SIC (figure from [39]) Figure 18: Three types of Hybrid III spines are available: a) curved spine, b) pedestrian straight spine, and c) FAA straight spine Figure 19: THOR spine (image from [43]) Figure 20: RID3D (image from [47]) Figure 21: BioRID II Spine (image from [48]) Figure 22: ADAM Spine (image from [50]) Figure 23: Spine criterion acceleration (image from [51]) Figure 24: Spine criterion force (image from [51]) iv

7 List of tables Table 1: Reported lumbar injury from vertical loading of PMHS specimens [13] Table 2: Mechanical properties determined from the FSU compression tests [14] Table 3: Summary of average FSU stiffness [17] Table 4: Summary of lumbar spine segment shear test results [17] Table 5: Summary of lumbar motion segment loads at failure [20] Table 6: Summary of results, combined dynamic bending and shear [22] and [23] Table 7: Summary of lumbar motion segment flexion-shear test results for Neumann et al[24] Table 8: Lumbar motion segment failure loads and their corresponding displacements [26] Table 9: BDRC model coefficients Table 10: Dynamic response limits, DR limit, for use in injury risk calculations (Equation 7) Table 11: Recommended military lumbar load limits [37] Table 12: Lumbar spine response to compressive loading Table 13: Lumbar spine response to shear loading Table 14: Lumbar spine response to flexion Table 15: Lumbar spine response combined flexion and shear loading v

8 1 Introduction Problem Definition In recent military conflicts in Iraq and Afghanistan, the incidence of lumbar spinal injuries sustained by mounted US military personnel from underbody blast (UBB) represents 26% of the 1819 reported incidents of spinal injuries or 145 cases. For the most part, the thoracolumbar junction from the T10 vertebra to the L3 vertebra was the most susceptible region of the spine to injury [1]. An increase in spinal injuries was noted following the introduction of vehicles designed to protect against underbody blast through the use of V-shaped hull design, energy absorbing seating systems and increased ground clearance. However, this increase may be more related to the change in the adversaries tactics as the conflict moved from Iraq to Afghanistan where buried improvised explosive devices (IED) became the weapon of choice. In Bernstock et al s [2] review of combat related spine injuries, it was noted that amongst Canadian Forces (CF) casualties reported at the Role 3 Multinational Medical Unit in Afghanistan there was an 8% incidence of spine injuries of which 10.4% were from IED blast, however, there was no mention as to the proportion of these that were sustained by mounted soldiers. From personal discussions with DRDC, the prevalence of lumbar spine injuries from underbody blast amongst CF personnel has been noted to result in debilitating long term effects on mounted CF servicemen and women. DRDC Valcartier is investigating the spinal injuries that are associated with exposure to UBB threats. Of interest is the effect of seating posture and increased inertial loads from body armour and torso borne equipment on the loads applied to the lumbar spine during an UBB incident. The energy absorbing seats that have been implemented in CF armoured vehicles do provide protection against UBB loading but the level of protection cannot be properly assessed using current test methods. The current Hybrid III anthropomorphic test device (ATD) used in experimental blast trials along with the Dynamic Response Index (DRI) for assessing lumbar spine injuries, are believed to be inadequate for assessing the various modes of dynamic loading and related injury risks. While the current method may be suitable for certain upright seating postures for which the ATD was designed, it is not amenable to the slouched posture often assumed by mounted crew members as it may not properly reproduce the loading for a humanlike seating posture. Furthermore, the DRI injury criterion only uses the vertical acceleration of the pelvis in its computation of injury severity which would underestimate the loading due to misalignment of the rotated pelvis accelerometer s measurement axis. In this orientation, injury is likely to involve bending and shear components in addition to the axial compressive loading of the spine. In their retrospective study of injured NATO and Afghan soldiers during Operation Enduring Freedom in Afghanistan, Ragel et al [3] found that flexion-distraction fractures represented 1.8% of all trauma admissions to the Craig Joint Theater Hospital and 38% of the reported thoracolumbar fractures. This type of fracture to the thoracolumbar spine results from hyperflexion of the spine about a fulcrum. In the case of the mounted soldier, their body armour and attached gear may act as the required fulcrum during vertical loading from the initial UBB launching the vehicle in the air or from the impact as the vehicle hits the ground. Possible loading mechanism that could result in flexion-distraction fractures are presented in Figure 1. Scenario A involves an unbelted soldier equipped with body worn gear subject to vertical loading who subsequently folds about the bulk of their gear as the body compresses under load. 1

9 Scenario B involves an unequipped soldier with lap restraint who flexes and collapses downward during vertical loading. Finally, Scenario C involves an unequipped soldier restrained by lap and a torso belt whose pelvis rocks forward and rotates underneath the lap belt as the legs are thrown upwards resulting in a reverse flexion of the thoracolumbar spine. These loading mechanisms may be further exacerbated with a poor initial seating posture. Figure 1: Possible loading mechanism that may result in fracture (figure from [3]). Although vertical loading is present in all three scenarios shown in Figure 1, compressive loading of the spine is not the underlying mechanism of lumbar injury but rather hyper-flexion which is not assessed by the Dynamic Response Index, the current injury assessment reference value used in full scale testing to determine spinal injury outcome following UBB. The work presented herein will support DRDC s investigation of lumbar spine injury and the development of a surrogate spine to realistically assess seating systems performance during UBB experiments where the influence of posture and inertial effects on potential injury mechanisms associated with non-ideal seating postures and not well known. 2

10 2 Effects of Seating Posture Blast attenuating seats attempt to use energy management mechanisms to reduce the forces from a UBB from being transmitted directly to the vehicle occupants. These attenuation mechanisms may include, dampers, padding materials, seat belts and alternative seat mounting methods. Regardless of a seat s construction, its effectiveness at reducing injury is partly predicated on the occupant being able to attain a proper seating posture. In the case of a soldier sitting in a military vehicle this may not always be possible. With the bulk and mass of a soldier s standard issue personal protective equipment, obtaining a correct seating posture may be difficult to achieve and often an out-of-position slouched seating posture may be assume. Generally, poor posture in one s daily activity (sitting, standing, and lifting) is detrimental to the load capacity and stability of the spine, whether during low level loads, as evident in chronic injuries, or with acute loads potentially leading to strains, extractions or fractures. There is limited data available pertaining to high rate vertical loading of the spine and associated lumbar injuries, particularly as they relate to seating posture. That being said, Liu et al [4] investigated the effects of spinal positioning of pilots during ejection using a mathematical model of the spine. In their review they cited earlier research that showed that 17% of ejection seat deployment resulted in spinal fracture of which 70% of the injuries were to the thoracolumbar spine. In their analysis they came to several conclusions related specifically to ejection seats and injury mechanisms, the most pertinent to the current UBB study are; (a) spinal alignment at the time of ejection determines the location and magnitude of stresses applied to the spine, and; (b) seat back restraints reduce the normal stresses and the use of a lumbar pad can reduce anterior stresses by placing the lumbar spine in extension. In Latham s [5] investigation of safe ejection speeds and accelerations, he noted that compressive loading to the spine can only be safely tolerated if flexion of the spine can be prevented. Otherwise, as documented in cases of inadvertent ejection-seat deployment in which the pilot was incorrectly positioned, that compression fractions to the T11 to L2 region were likely. It was found that correct spinal alignment could be achieved by means of a 3.8 cm (1.5 in) lumbar pad on the seat back. Adams et al [6] conducted testing on excised lumbar sections to evaluate the effects of posture on the compressive strength of the specimen. The objective was to establish a range of flexion and extension in which the lumbar spine is most resistant to compressive loads. Nineteen lumbar spines were dissected into 29 motion segments consisting of two vertebrae and the intervening disks and ligaments. The vertebrae were potted in dental plaster and secured to two endplates that could be loaded in a computer controlled hydraulic materials testing machine. The endplates could be angled relative to one another to simulate compression, bending and shear that may act on a motion segment. The lower endplate was fixed to the hydraulic ram whereas rollers on the upper endplate were free to roll on a bearing plate attached to a load cell. The height of the rear roller was adjustable such that the various bending moments could be induced for a given compressive load. The motion segments were loaded until the elastic limit was reached as indicated by a reduction in slope of the moment versus flexion angle output graph. The results of their tests suggest that the lumbar spine is best able to withstand compressive loading when the motion segment is flexed within 0% to 75% of its flexion range and that the optimum stress distribution in the disk occurs at 50% of maximum flexion, this appears to contradict Latham s 3

11 conclusion that compressive loading can only be tolerated if flexion of the spine can be prevented. Preload of the lumbar spine can be introduced by the mass of body worn personal protective equipment or other equipment acting to apply a compressive load on the intervertebral discs. Additionally, postural changes that result in non-neutral spine positioning can also introduce compressive loads from the muscle s contraction required to maintain a stable spine. These postural changes could be the result of slouching or raised leg position when seated. Tension in the musculature acts only to stabilize the spine and in so doing increases the compressive loads on the intervertebral discs. However, similarly to a neutral spine position the muscles cannot react fast enough to effect body kinematics during an UBB. Janevic et al [7] conducted experiments on thirteen motion segments that were mounted in a test apparatus that used a system of cables and pneumatic cylinder to apply the loads and preloads to the specimen. The preload was kept constant while various bending moments, shear loads and one torsion load were applied. Four incrementally increasing loads of 40 N were applied up to a maximum of 160 N. Similarly, an incremental moment of 4 Nm was applied up to a maximum 16.0 Nm. The sequence of test loads were introduced while a compressive preload of 0 N, 2200 N or 4400 N was applied. The results of the testing indicated that the stiffness of the lumbar spine in bending, shear and torsion would increase when compressive loads are applied. Figure 2 and Figure 3 show the decrease in displacement (mm or deg) for the applied test loads of 80 N or 8 Nm for shear and flexion respectively. Figure 2: Mean translational displacements of lumbar spine segments under preload conditions. 4

12 Figure 3: Mean rotational displacements of lumbar spine segments under preload conditions. 5

13 3 Upper Body Mass Effects On Lumbar Injury To be effective, blast attenuating seats must be tuned to the mass of the occupant. If the energy absorption mechanism is too stiff for a given occupant then the seat may not use its maximum displacement range or ride-down thereby increasing the loads on the occupant. Conversely, if the seat is not stiff enough it may bottom out resulting in very high loads being transmitted. To complicate matters, a military vehicle occupant may be wearing a significant amount of extra mass due to the protective equipment or other gear that may be worn on their person, typically on the upper torso degrading the performance of the attenuation system. Iluk [8][9] used LS-Dyna to evaluate blast loading on three seating configurations. The first seat configuration employed a constant damping force set for the mass of the ATD used in a typical STANAG 4569 test. In the second configuration, the constant damping force was proportional to the real mass of the ATD and the third configuration the damping was proportional to the measured mass of the ATD including any protective gear worn. The ratio of the actual strain in the spine to the strain of the base ATD without additional mass was used to compare risk based on DRIz calculations. For similar loading, a 40% increase in spine loads was observed for the ATD with the increased torso borne mass while at the same time the calculated DRIz decreased due to the lower accelerations that would be imparted to a heavier occupant. An alternative method for computing DRIz that accounts for additional body borne mass was proposed such that the total occupant mass should be considered in the calculation of DRIz. The rationale being a seat designed for a 50 th percentile occupant may not provide the correct damping for a 5 th percentile occupant or for the 50 th percentile subject wearing heavy ballistic armour. In a helicopter crash, aircrew are subjected to vertical loading when hitting the ground and it has been suggested that this loading is similar to that experienced in a surface vehicle hitting the ground after been launched vertically from the force of a UBB event. The resulting loads on an aircrew s spine may be exacerbated by the extra mass of the body borne equipment. This may include body armour and primary survival gear and other gear such as ammunition, radio, knife and flashlight that can add between 5 kg to 30 kg of extra mass. Using a finite element model of a 50 th percentile Hybrid III, Aggromito et al [10] further evaluated the effects that mass distribution and placement of the extra items have on the resulting loads and injuries. Aggromito also investigated whether the mass increase causes the seats to be overloaded in a crash. As indicated, the seats are designed to absorb vertical loading by collapsing, however, if the total mass on the seat exceeds its design specifications the seat may bottom out resulting in much higher vertical loads on the occupant. A Hybrid III ATD with a curved spine was used in their testing although the straight spine variant would have been preferred but was not available. The findings as they relate to lumbar loads are summarized here: lumbar loads increased when seat bottoming occurred, the sooner bottoming occurred the larger the lumbar loads, equipment placed on the sides of the body produced lower lumbar loads compared to the equipment placed on the front of the torso, and; lumbar loads increased with equipment bulk which caused the body to move forward, the more offset from the body the larger the lumbar loads. 6

14 Zhang et al [11] used a high-fidelity finite element model of the pelvis and lumbar spine to evaluate the potential of lumbar injuries as it relates to the loading severity during the initial stages of high-rate vertical loading. The effects of added torso borne mass on injury outcome were also assessed. The lumbar FE model used was developed by Johns Hopkins University Applied Physics Laboratory (JHU/APL) and is based on geometry obtained from the Visible Human Project 1 and scaled to the dimensions of the average US male. The input to the pelvis and lumbar spine model was determined from FE simulations with a full body Hybrid III dummy and applied to the ischial tuberosity. The torso and torso borne equipment were represented by a couple of masses connected to the upper end of the lumbar spine. The results from the FE simulations suggested that at high rate loading the added torso borne mass had little effect on the injury outcome since under severe pelvis acceleration lumbar spine failure would occur regardless. The extra mass may have had an effect during later stages of the loading event which were not included in the model run times. It was noted that for lower severity pelvis loading conditions, the added torso borne mass might have an effect on the maximum lumbar loads. 1 For more information on the Visible Human Project please see: 7

15 4 Spine Response to Loading Currently, the Hybrid III antropomorphic test dummy is used in the assessment of occupant injury during UBB full-scale test programs. However, its lumbar spine was designed for forward flexion that is common in automotive frontal crashes but it was not intended for the vertical loading mechanisms of an UBB environment. Additionally, because of the Hybrid III lumbar spine s stiffness and curvature, it is difficult to position the dummy in the blast attenuation seating systems or to have it assume a posture that is less than ideal (i.e. slouched). A stated objective of the current development program is to design a new functional spine unit (FSU) that could replace the current Hybrid III spine having greater biofidelity and ability to assume more humanlike postures. In order to accomplish this, specifications for the spine s stiffness and range of motion must be developed. The results of a literature review seeking information related to the mechanical characteristics of the lumbar spine are presented below. Most research reported in the literature considered ideal loading conditions and therefore necessitates additional fundamental research to develop lumbar tolerances for out-of-position loading. Regardless, the data that is available can potentially be used in developing a preliminary definition of a surrogate spine s response to UBB loading. 4.1 Response to Compression Loading Stemper et al [12][13] simulated the vertical loading on the spine during the catapult phase of an aircraft ejection seat and during a helicopter crash to investigate the fracture patterns from these high-rate loading events. A drop tower arrangement was used to apply vertical loading to post mortem human subject (PMHS) spine segments (T12-L5). The upper end of the spine was fixed to an upper plate such that the L2-L3 intervertebral disk was horizontal. The lower end of the spine was attached to a bottom plate via a 6 axis load cell. An accelerometer mounted to the bottom plate measured the input into the system. The upper torso mass was simulated with a 32 kg mass attached to the upper plate. The specimens were pre-flexed with a 5 Nm moment. Both the upper and lower plates were independently attached to a rail system which guided the drop assembly in a vertical orientation for impact with a foam pad. Pad materials were selected to simulate either the loading experienced during the catapult phase of an emergency ejection or during a helicopter crash. Compression fractures occurring during the initial loading phase were identified as the primary mechanism of injury. The peak forces and moments are summarized in Table 1 but there was no indication of the magnitude at which the injuries occurred. Stemper recognized that the results are likely dependent on other factors such as column length, curvature, bone quality, spine orientation and state of flexion or extension when loading is applied, but these variables were not investigated further. 8

16 Table 1: Reported lumbar injury from vertical loading of PMHS specimens [13]. Type Peak Axial Force (N) Peak Moment (Nm) Injury Description Ejection Catapult L1- Anterior cortex fracture with endplate fracture Ejection Catapult L1- Burst fracture with retropulsion into spinal canal Helicopter Crash L1 and L2- Anterior compression Helicopter Crash L3- Vertical cortical fracture with laminar fracture L4- Mild compression fracture of the anterior cortex Helicopter Crash L2- Anterior body fracture including cranial endplate L2- Vertical cortical fracture of posterior wall L3- Burst fracture with retropulsion into spinal canal Yoganandan et al [14] conducted compression tests on 9 FSUs from 6 different donors. In four of the specimens, the intervertebral disks of the FSU showed signs of degeneration. The specimens were potted and installed in a servo-hydraulic material test system. A load cell in the test apparatus loading head measured the axial compressive load that was applied at a rate of 2.54 mm/s normally to the mid-plane of the FSU s intervertebral disk. A typical force deflection curve is shown in Figure 4 for both a normal and a degenerated intervertebral joint. The measured mechanical properties are summarized in Table 2. Figure 4: Compressive force deflection curve for; (a) normal, and; (b) degenerated intervertebral disk loaded in compression (figure from [14]). 9

17 Table 2: Mechanical properties determined from the FSU compression tests [14]. Overall Stiffness Load at Failure Ultimate Load 2850 ± 293 N/mm 9.02 ± 1.08 kn ± 1.42 kn Years later, Yoganandan et al [15] investigated the mechanism of compression related injuries of the thoracolumbar spine seen during automotive frontal crashes. Laboratory tests comprising vertical impact tests on spinal columns obtained from unembalmed post mortem human subject (PMHS) were conducted. Similar to Stemper, the lower end of the spinal section was potted and fixed to a metal plate through a load cell. The upper end was similarly potted and fixed to an upper metal plate with 5 Nm of induced flexion. The entire assembly was dropped in guided free fall on a custom drop tower with the bottom plate striking an impact pad, simulating the loading from a seat into the pelvis and lumbar spine of a vehicle occupant. An accelerometer in the lower steel plate measured the input into the system. Three drops of increasing energy levels were conducted for each specimen; the first two drop energies were non-injurious, while the third was sufficient to cause injury. The extent of injury was assessed using pre and post-test radiographs and CT scans. The authors recognized that repeated impacts to the same test specimen may decrease its tolerance levels during subsequent impacts, therefore, the results obtained are considered to be conservative estimates of the failure loads. A logistic regression analysis indicated that a peak compressive force of 3.7 kn is associated with a 50% probability of injury in the thoracolumbar spine. The authors speculated that if only the lumbar spine was included in the test that the load required to cause a 50% probability of injury would be greater than 3.7 kn, so again the results are conservative. The logist curves for probability of injury obtained are presented in Figure 5. Figure 5: Probability curves for spinal injury due to compressive loading (figure from [15]). Kemper et al [16] investigated the effect of loading rate on the compressive stiffness of lumbar intervertebral disks. Thirty-three (33) compression tests were conducted on 11 FSU obtained 10

18 from six fresh frozen human lumbar spines. The cranial vertebral body of the FSUs were potted such that the mid-plane of the disk was parallel to the potting cup and attached to the test apparatus. The caudal vertebra was lowered into the second potting cup which prevented any moment from being introduced. The test apparatus comprised of a hydraulic material test system with a five-axis load cell for measuring the reactionary loads and a single axis load cell for measuring the applied impact load. Accelerometers were also placed on the impactor and the impact plate. The FSUs were preconditioned with 10 cycles of a 0.5 mm displacement at 0.1 m/s and then preloaded to N followed by a 0.5 mm and a 1.0 mm dynamic step inputs loaded at a rate of 0.1 m/s and 0.2 m/s respectively. The preload of N was reconfirmed and the FSUs were then loaded to failure at a rate of 1 m/s. The measured stiffness was found to be independent of the vertebral level from which the FSU had been harvested but it increased with the rate at which the specimen was loaded, as shown in Figure 6. Figure 6: Compressive stiffness of intervertebral disks by lumbar level and loading rate [16]. The average stiffness for each loading rate is summarized in Table 3. established using the measured disk heights. The strain rates were Table 3: Summary of average FSU stiffness [17]. Loading Rate (m/s) Average Stiffness (N/mm) Strain Rate (s-1) ± ± ± ± ± ±

19 Combining his results with previous published data, Kemper proposed the relationship between stiffness and strain rate shown in Equation (1) and the effective stiffness for the full lumbar spine is given by Equation (2). k = ε (1) k eff = 1 N [ 1 1 k ] N ; N = 5 (2) 4.2 Response to Shear Loading The shear strength of the lumbar spine as it relates to bone mineral density (BMD) and disk height reduction was investigated by Skrzypiec et al [17] in laboratory experiments using 29 L2- L3 motion segments comprised of the intervertebral disk, adjacent vertebrae, apophyseal joint and ligaments. The motion segments were categorized into three groups: young-no-creep, youngcreep and finally old-no-creep. Creep is the reduction in disk height due to a sustained compressive load on the disk. For these tests creep was introduce by applying a 1000 N load for one hour prior to testing. The vertebrae were positioned in metal cups and embedded in resin. The cups were mounted to a servo-hydraulic testing machine such that the axis of the spine was horizontal with the anterior side facing up. The caudal vertebra was fixed to the testing machine via a 6-axis load cell and a dead weight and cable/pulley system applied a 500 N compressive load to the joint while fore-aft shear was introduced through a 15 mm displacement of the cranial vertebrae. Throughout testing, the motion segment was immersed in a Ringer solution to maintain its hydration. The force deflection during the initial stages of shear displacement were found to be somewhat linear in both the anterior and posterior directions, however, the stiffness was higher in the anterior direction which is likely due to the overlapping geometry of the apophyseal joint. Not all the segments failed within the 15 mm displacement of the cranial vertebra but failures did typically demonstrate load-displacement curves that reached a plateau rather than failing in a brittle manner. It was further noted that the motion segment exhibited an increase in height as the shear displacement was applied. A summary of the results is presented in Table 4 below and sample load-displacement curves are presented in Figure 7. Table 4: Summary of lumbar spine segment shear test results [17]. Group Stiffness Stiffness after creep Yield Peak Anterior (N/mm) Posterior (N/mm) Anterior (N/mm) Posterior (N/mm) Force (kn) Disp. (mm) Force (kn) Disp. (mm) Young-No-Creep Young-Creep Old-No-Creep

20 Figure 7: Typical shear results for a L2-L3 motion segment (figure from [17]). 4.3 Response to Flexion-Shear Loading Schultz et al [18] tested 42 motion segments obtained from 24 fresh lumbar spine sections. The cranial and caudal sections were potted in epoxy resin and secured in a cylindrical test fixture. The caudal fixture was clamped to the test machine base while the cranial section was unconstrained. Through a series of wire attachments to the specimen s cranial support tube flexion, extension and lateral moments could be applied to the motion segment. A reconfiguration of the wire connections to the specimen s test fixture allowed for torsional loading to be applied. Bending and torsion loads up to 20 Nm and shear loads up to 205 N were applied to the cranial support fixture using the wire loading system. A constant 400 N compressive load, approximating the weight of the body above L3, was applied for all the tests. The mean responses for the testing of the different loading modes were presented graphically and are shown in Figure 8. The tests were repeated with the posterior elements of the motion segments removed. 13

21 Figure 8: Mean flexion, extension, lateral bending and torsion response of motion segments tested by Schultz et al [18]. Dashed lines are tests with the posterior elements removed. In follow-on testing to that conducted by Schultz [18], Berkson et al [19] conducted compression and shear tests on the same motion segment specimens using the same test machine to apply shear and compressive loads. The results of these additional tests are shown graphically in Figure 9. 14

22 Figure 9: Mean compression, lateral, anterior and posterior shear response of motion segments tested. Dashed lines are tests with the posterior elements removed (figure from [19]). Osvalder et al [20] conducted laboratory experiments to assess the ultimate strength of the lumbar spine in flexion. Sixteen motion segments comprising an intervertebral disk and the adjacent vertebra were subjected to an increasing static bending and shearing load up until the point of complete failure. The motion segment specimens were extracted from PMHS lumbar spines (T12-L5) and consisted of L1-L2 vertebra, L2-L3 vertebra or L3-L4 vertebra. For the testing, the cranial and caudal vertebrae were potted in metal cups. The caudal cup was rigidly attached to a force plate for measuring the reaction loads from a moment applied to the cranial vertebra by a vertical lever arm affixed to the upper mounting cup. Regardless of the composition of the FSU, complete failure occurred at an applied moment of 156±11 Nm with a corresponding shear of 620±53 N and there was no correlation between the lumbar spine levels tested. None of the specimens failed below 135 Nm in bending and 520 N in shear. Just prior to failure, the tension in the rear portion of the motion segment was 2.8±0.2 kn. A strong correlation between the bone mineral density (BMD) of the adjacent vertebra and the ultimate strength was noted, as was also noted in Neumann et al s [21] analysis of the data. Moment versus angle-of-rotation for the cranial vertebra is shown in Figure 10. The force and moment results are summarized in Table 5. 15

23 Figure 10: Flexion and versus bending moment at failure of lumbar motion segments (figure from [20]). Table 5: Summary of lumbar motion segment loads at failure [20]. Spine Level At Failure Maximum Bending Moment (Nm) Shear Force (N) Tensile Force posterior structure (N) Flexion (deg) Horizontal Displacement (mm) L1-L2 154 ± ± ± ± ± 1.4 L2-L3 153 ± ± ± ± ± 1.5 L3-L4 161 ± ± ± ± ± 2.4 Osvalder et al [22] conducted additional experiments on 20 motion segments to look at their response to combined dynamic bending and shear. L1-L2 or L3-L4 motions segments were mounted in a test fixture with the caudal end affixed to a load and rigidly attached to the ground. A 12 kg mass applied a preload on the end of the motion segment. A combined shear and bending moment was applied with a 20 kg padded pendulum that struck a vertically mounted leaver arm that was rigidly attached to the cranial vertebra. Either a moderate or severe pendulum impact, as defined by an accelerometer on the pendulum, was applied to the specimens. A pendulum acceleration of 5 G with a rise time of 30 ms and duration of 150 ms was considered moderate whereas an acceleration of 12 G with a rise time of 15 ms and duration of 250 ms was considered severe. The results of the testing are presented in Table 6 below. Using the same test methodology, Osvalder et al [23] conducted a similar test series on 48 additional motion segments. However, in addition to the moderate and severe test conditions a medium condition defined by an 8 G pendulum acceleration pulse was also added. The data for all the tests from both test programs are summarized in Table 6 below. The values presented are the maximum before a noticeable dip in the output was observed. 16

24 Table 6: Summary of results, combined dynamic bending and shear [22] and [23]. Impact Severity Peak Accel. (G) Rise Time (ms) Moment (Nm) Shear Force (N) Flexion (deg) Horizontal Displacement (mm) Vertical Displacement (mm) Moderate ± ± ± ± ± 1.4 Moderate ± ± 73 NA NA NA Medium ± ± ± ± ± 1.1 Medium ± ± ± ± ± 1.8 Severe ± ± ± ± ± 2.4 Severe ± ± ± ± ± 2.5 There were indications of injury onset in all the tests that was evident by a distinct dip in the measured force and moment response as seen in a typical bending moment response shown in Figure 11. Figure 11: Typical bending moment versus time response for severe load with a 15 ms and a 5 ms pulse rise time (figure from [23]). Neumann et al [24] determined the injury threshold of 11 motion segments loaded in combined flexion and shear. Four spines were sectioned into motion segments comprising T12-L1, L2-L3 and L4-L5. The specimens were potted and loaded and unloaded in 20 N increments until the first signs of injury were detected as evident by a dip of 50 N or 15 Nm in the force deflection or moment deflection responses respectively. The load was applied in the anterior direction to a vertical lever arm affixed to the cranial vertebra of a vertically mounted motion segment. Permanent deformation was recorded at an applied moment of 121 Nm (SD = 10 Nm) or 486 N shear force (SD = 38 N). Pertinent results from Neumann s experiments are summarized in Table 7. 17

25 Table 7: Summary of lumbar motion segment flexion-shear test results for Neumann et al[24]. Parameter Bending Moment (Nm) 121 (10) Shear Force (N) 486 (38) Horizontal Displacement (mm) 7.5 (1.4) Vertical Displacement (mm) 16.9 (1.4) Flexion Angle (deg) 15.8 (0.9) Bending Stiffness (Nm/deg) 7.5 (1.6) Shear Stiffness (N/mm) 71 (7) Measured result (standard deviation) Adams et al [25] conducted similar experiments to Osvalder with the exception that a compression creep test was performed on 15 PMHS specimens to ensure the water content of the disc lay within the physiological range. This was accomplished by applying a compressive force of between 1000 N and 2000 N depending on the specimen s age for a period of 2 and 3 hours. This procedure reduced the overall height of the motion segments by 1.09±0.4 mm. The motion segments were then flexed to failure. The elastic limit of the specimens was reached at 72.8±18.1 Nm with a rotation angle of 14.8±3.3. Whereas previous failure of the motion segment was identified by overt indications in the Osvalder experiments, Adams identified failure by a slight reduction in the loading curve indicating that the elastic limit of flexion had been reached and that at this stage of flexion only the supraspinous/interspinous ligaments are damaged which would not have been noticed in the previous experiments. Moment versus angle-ofrotation for a typical test is shown in Figure 12. Figure 12: Moment deflection curve up to the elastic limit of the FSU (figure from [25]). In her master s thesis, Belwadi [26] investigated combined anterior shear and flexion of the lumbar spine that may occur in a frontal automotive crash in which knee bolsters are deployed. Knee bolsters could restrict forward motion of the pelvis allowing for a greater relative motion between the upper and lower torso. 18

26 The response of 7 full cadaveric lumbar spines (T12-Sacrum) to sub-injurious tension-compression, flexion-extension, shear, and combined loading was measured in a spine testing machine that could apply a fixed displacement and/or rotation to the cranial end of the test specimen. The full spine specimens were then sectioned through the L3 vertebral to obtain 11 motion segments comprising three vertebra and two discs. Each motion segment was also tested similarly to the full sections. Failure in the combined shear-flexion mode was noted in eight of the 155 total tests that were performed. The failure forces and moments for these tests are summarized in Table 8. Table 8: Lumbar motion segment failure loads and their corresponding displacements [26]. Motion Segment No Shear Flexion Force (N) Displacement (mm) Moment (Nm) Angle (deg) Sacrum to L ± ± ± ± 1.4 L2 to T ± ± ± ± 2.3 Average ± ± ± ± Range of Motion There is limited information in the literature specific to lumbar range of motion that would be pertinent to the development of an ATD spine. There is a plethora related to range of motion for various daily activities and chronic back pain, however, these were specific to the modality being investigated. A comprehensive review of spine mobility was reported by Nyquist and King [27]. Nyquist and Merton [28] performed experiments on six male volunteers to establish static range of motion from T8 to the pelvis. Full range of motion with straight legs was 120 but increased to 143 when legs were flexed 90. Dynamic range of motion was determined by Cheng et al. [29] using a test sled for accelerations up to 8 G. Range of Motion (ROM) for T12/Pelvis was determined for both male and female subjects, and in a tensed and relaxed condition. Angles (± 1 SD) of 22.7±6.8 to 28.0±4.3 degrees were observed. Demetropoulos et. al. [30] looked at the flexion and extension properties of cadaveric lumbar specimens, both with and without simulated muscle tension. Without muscle tension, the flexion results were quite variable, however, when simulated muscle tension was applied, the results became much more coherent (96.5 ±1.76). For the most recent frontal automotive ATD development program, the THOR dummy, it was determined that the existing information was insufficient to specify any spine kinematic response requirements [31]. 19

27 5 Existing Injury Criteria 5.1 Eiband Injury Tolerance Curve for Vertical Acceleration The Eiband injury tolerance curve was proposed for assessing the risk of injury from vertical impacts [32][33]. It suggested that under vertical loading conditions, human volunteers tolerated acceleration levels of 10 G for 0.1 s and 15 G for 0.05 s. Adding data obtained from animal tests the criteria was expanded to exposure limits for moderately and severely injured occupants. The limitation of the criterion is that it did not distinguish amongst body regions and it was not sensitive to pulse shape and or duration. The Eiband tolerance curve is shown in Figure 13. Figure 13: Eiband injury tolerance curve for vertical acceleration (figure from [32]). 5.2 Dynamic Response Index The Dynamic Response Index (DRI) [33] is used as a lumbar spine injury assessment reference value for blast tests of vehicle and seating systems. The calculation for DRI is based on the response of a lumped mass model representation of the spinal column (see Figure 14). The criteria DRIz denotes loading in the vertical direction. The input into the model was originally the vertical acceleration of the seat pan of an aircrew ejection seat but its use has been modified for blast attenuation using the vertical acceleration of the ATD s pelvis during the experimental trials. The maximum value of the deflection response is the value of DRI which has been correlated to experimental and operational injury data (see Figure 15). 20

28 Figure 14: DRI lumped mass model of the spine (figure from [33]). Figure 15: DRI versus spinal injury rate (figure from [32]). The equation of motion for DRI s lumped mass model is: Where: Ζ (t) = δ + 2 ζ ω n δ + ω n 2 δ (3) Ζ (t) is the pelvis acceleration in the vertical direction δ = ξ 1 ξ 2 (when delta > 0) is the compression of the system ζ = c 2 m ω n is the damping coefficient (0.224) ω n = k is the circular frequency (52.9 rad/s) m 21

29 DRI is calculated with the maximum compression δ max, ω n 2 and the acceleration of gravity g is (9.81 m/s): DRI z = ω n 2 δ max g (4) A DRIz value of 17.7 is currently used as the pass-fail criteria in the AEP-55 test procedures [34]. Spurrier et al [35] suggests that the DRI criterion is inappropriate for use in assessing the probability of spine injury in vehicle blast scenarios. A comparison of injuries locations amongst aircrew during ejection, for which DRI was developed, and those of vehicle occupants subjected to UBB showed that the location of injury varied significantly between the two events. Spine injuries from aircraft ejection were primarily to the thoracic spine whereas injuries from UBB we predominant in the lumbar spine, suggesting the mechanisms of injury are very different. Furthermore, DRI is intended for assessing the probability of injury to vertical loading and is not specific to the lumbar spine. In an UBB, the deformation of the vehicle floor pan may push up on the occupant legs thereby rotating the pelvis and lumbar spine such that they are no longer aligned with the direction of loading. The DRI would then underestimate the severity of the loading. 5.3 Multi-axial Dynamic Response Criteria The Multi-axial Dynamic Response Criteria is referred to as the Brinkley Dynamic Response Criteria (BDRC) by the National Aeronautics and Space Administration (NASA) [36]. It is similar to the computation of the DRI except that the BDRC is based on the response of three lumped mass models representing the three orthogonal axes with the origin located at the critical point and is used to calculate a general whole body injury risk. For a seated occupant the critical point is fixed relative to the seat coordinate system and shown in Figure 16. Figure 16: Critical point for computing BDRC (figure from [36]). 22

30 The equation for the BDRC lumped mass models, for each of the three orthogonal directions, is given by Equation 5. Where: A(t) = δ (t) + 2 ζ ω n δ (t) + ω n 2 δ(t) (5) δ (t) is the occupants acceleration in the inertial frame δ (t) is the occupants relative velocity with respect to the critical point δ(t) is the displacement of the occupant s body with respect to the critical point ζ the damping coefficient ratio ω n undamped natural frequency of the dynamic A(t) is the measured acceleration, per axis, at the critical point. Rotational accelerations must be considered when computing the linear accelerations at the critical point The coefficients to be used in the computation of the BDRC for each of the three orthogonal axes are presented in Table 9. Table 9: BDRC model coefficients. Coefficient X displacement direction Y displacement direction Z displacement direction X<0 X>0 Y<0 Y>0 Z<0 Z >0 ω n ζ For each axis, the dynamic response is calculated according to Equation (6) and the risk of injury (β) is obtained from Equation (7). The dynamic response limits for each axis are presented in Table 10. The injury risk (β) is first calculated using the DR limit for the low risk and if it is less than 1 then the value represents the risk of injury. If it is greater than 1 then the injury risk is recalculated using the next higher DR limit. DR = ω n 2 δ(t) g (6) β(t) = ( DR 2 x(t) lim DR ) + ( DR 2 y(t) lim x DR ) + ( DR 2 z(t) lim y DR ) z (7) 23

31 Table 10: Dynamic response limits, DR limit, for use in injury risk calculations (Equation 7). Axis X Y Z Displacement direction Negative Positive Negative Positive Negative Positive Low (<0.5%) -28 DR x < 0 0 DR x < DR y < 0 0 DR y < DR z < 0 0 DR z < 15.2 Medium (0.5% to 5%) -35 DR x < DR x < DR y < DR y < DR z < DR z < 18.0 High (5% to 50%) -46 DR x < DR x < DR y < DR y < DR z < DR z < 22.8 Although the risk of injury can be computed from easily measured accelerations, the over simplification of the human-vehicle interactions to solely an acceleration input model may not assess the injury from other potential injurious causes. Additionally, the limits used in the calculation of BDRC are derived from a specific ejection seat test setup and their use for different loading conditions may require additional verification or validation. 5.4 Lumbar Load Criterion The Crew Systems Crash Protective Handbook [37] specifies lumbar load criteria for assessing seat performance in helicopter crashes. The assumption is that the severity of injury is directly related to the magnitude of the loads applied in a crash. The lumbar load criterion is based on a FAA criterion for civil aviation seats that specifies a maximum load limit of 1500 lbs (6672 N) measured between the pelvis and the lumbar spine of a 50 th percentile crash test dummy for vertical loads in line with the spinal column. This relates to a 9% risk of a detectable spinal injury occurring. The Naval Air Warfare Center adopted the load criteria approach but used scaled threshold load values for different occupant sizes, as presented in Table 11. These values were obtained from horizontal accelerator sled tests with Hybrid III 5 th, 50 th and 95 th percentile ATDs with the straight spine and articulating hips of the aerospace variant of the ATD [38]. It was found that the use of the ATD s abdominal insert had a significant effect on the measured loads as it provided an alternate load path to transmit loads from the pelvis to the upper torso. Table 11: Recommended military lumbar load limits [37]. Occupant Size Lumbar Load Tolerance (lbs) 5 th percentile female 1281 (5698 N) 50 th percentile female 1610 (7162 N) 50 th percentile male 2065 (9186 N) 95 th percentile male 2534 (11272 N) As with the DRI results, the lumbar loads can be affected by the orientation of the seat relative to the loading direction and by the positioning of the ATD in the seat [33]. 24

32 5.5 Spine Injury Criterion Dosquet et al [39] proposed a spine injury criterion (SIC) that is based on the effective impulse to the spine and is calculated by integrating the applied force to the spine over time. In an analysis of over 200 tests on a vertical drop tower where seated Hybrid III manikins were dropped in free fall, a correlation between SIC and DRI was obtained. It was also determined that the best correlation was obtained if the difference between the limits of integration was 40 ms with the use of an exponent of 1.2. The SIC equation is: SIC = sup t1, t 2 {( 1 t 2 t 1 t 2 Fdt) 1.2 (t t 1 2 t 1 )} with t 2 t 2 = 40 ms (8) The correlation between SIC and DRI as seen in Figure 17 is: DRI = SIC (9) Figure 17: Correlation between DRI and SIC (figure from [39]). 25

33 6 Anthropomorphic Test Devices There are several anthropomorphic test devices (ATD) that are available commercially and their suitability for UBB injury assessment is briefly discussed in the following sections. 6.1 Frontal Impact Dummies Hybrid III The Hybrid III ATD was developed by General Motors in the 1970 s to address the need for a more biofidelic crash test dummy than was available at the time. It was adopted into the United States Federal Motor Vehicle Safety Standards in 1986, and has become the worldwide standard for use in frontal impact crashworthiness testing. Due to its extensive use, and availability, the Hybrid III is also used in many other non-automotive applications where the use of an ATD is required (e.g. sports equipment testing, military testing), including UBB testing. The Hybrid III s biofidelity in many of these applications has not been demonstrated. The lumbar spine of the Hybrid III is a polyacrylate elastomer member which is rigidly attached to the dummies thoracic spine and pelvis. Two steel cables run through the mid-coronal plane of the lumbar spine, providing side support while permitting fore-aft flex. Three versions of the Hybrid III lumbar spine are produced as shown in Figure 18. The standard version is curved to permit the dummy to be accurately placed in a seated posture within an automotive vehicle. A straight version is also available to place the dummy in an erect stance for use in pedestrian impact testing, as well as a Hybrid II based straight spine version designed for aircraft seat testing which is incorporated into the FAA-Hybrid III dummy. a) b) c) Figure 18: Three types of Hybrid III spines are available: a) curved spine, b) pedestrian straight spine, and c) FAA straight spine. 26

34 The Hybrid III instrumentation includes: six-axis lower lumbar spine load cell, and; tri-axial pelvis acceleration. Injury risk functions for the spine when using the Hybrid III have centered on the use of DRIz. Two risk curves appear in the literature, one created from operational data and another from lab data [40]. The lab data presents a more conservative curve, so this has been used to calculate an Injury Assessment Reference Value (IARV) of 17.7 (10% risk of AIS2+) [40][41]. ATD s like the Hybrid III are instrumented to directly measure lumbar force. Compressive Lumbar Load (LL) uses the load cell input to develop Peak LL (Compressive) criterion. Attempts to create a criterion for LL have produced values of 5598 N to 6675 N by correlating compression to DRIz, or by scaling neck IARV data [40][42] THOR-M The THOR (Test device for Human Occupant Restraint) dummy was developed by the US Department of Transportation, National Highway Traffic Administration (NHTSA) to improve on the biofidelity of the Hybrid III ATD 2. It includes extensive instrumentation to assess the performance of vehicle safety systems designed to mitigate injuries in frontal and frontal oblique crashes. It is not currently used in the regulation of automotive safety systems. The lumbar spine is comprised predominantly of metal components, however there is a short (76 mm) butyl rubber flex joint to provide some lumbar compliance. This flex joint, like the Hybrid III, has two steel cables running internally on the mid-coronal plane to provide lateral stability. The lumbar spine is rigidly attached to the pelvis but an adjustable joint is provided at the thoracolumbar joint to aid dummy positioning. The Thor spine is depicted in Figure 19. Figure 19: THOR spine (image from [43])

35 THOR s relevant instrumentation includes: 5 axis lower thoracic (T12) load cell, tri-axial lower thoracic (T12) acceleration, and; tri-axial pelvis acceleration. No injury reference values specific to the THOR were found. The limited availability of the dummy is reflected in its restricted use in the field. FE modeling using the THOR was noted, with Lumbar Loading Criterion being used, and the values being the same as for the Hybrid III [44]. 6.2 Rear Impact Dummies RID3D The RID3D is a rear impact dummy developed in the early 2000 s as an improvement on the RID2 ATD. It is based on the THOR frontal impact dummy, having a THOR thorax, however, the lumbar spine is a straight cylindrical rubber component adopted from the EUROSID-1 dummy [45] (Figure 20). It has been validated for rear impact scenarios up to 60 km/h and 15 G [46] but no validation work has focused on impact biofidelity specific to the lumbar spine. Figure 20: RID3D (image from [47]). The RID3D has the same adjustment at the thoracolumbar connection as the THOR dummy. A group of researchers used the RID3D while testing lifeboats used on oil rig platforms. These boats are dropped from a height of 30 m, and in one case, the dummy is placed in a supine position with the legs crouched. They noted that the RID3D was the rear impact dummy with enough lumbar adjustment and flexibility to permit the correct seating position [46]. 28

36 The RID3D instrumentation includes: 6 axis lower thoracic (T12-L1) load cell, tri-axial lower thoracic (T12-L1) acceleration, and; tri-axial pelvic acceleration No lumbar injury assessment risk values are available for the RID3D. Rear impact dummy IARV work is concentrated on cervical spine injuries. A.1.1 BioRID II The BioRID II dummy was developed at Chalmers University through the late 1990 s into the 2000 s. The key distinguishing feature of the BioRID II is the spine which consists of 24 individual articulating components representing the human vertebrae (see Figure 21). Rubber bumpers are glued to the spine segments. Motion of the spine is constrained to the mid-sagittal plane. Figure 21: BioRID II Spine (image from [48]). The researchers who used the RID3D to assess drop launched life boats also considered the BioRID II, however, it did not have the same degree of lumbar flexibility and adjustment. Depending on the seating configuration, the BioRID II may offer better conformance to the seat [49]. 29