Engineering approaches to understanding mechanisms of spinal column injury leading to spinal cord injury

Clinical Biomechanics xxx (xxxx) xxx–xxx

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Clinical Biomechanics journal homepage: www.elsevier.com/locate/clinbiomech

Review

Engineering approaches to understanding mechanisms of spinal column injury leading to spinal cord injury Claire F. Jonesa,b, Elizabeth C. Clarkec,

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a

Spinal Research Group, Centre for Orthopaedics and Trauma Research, Adelaide Medical School, The University of Adelaide, Australia School of Mechanical Engineering, The University of Adelaide, Australia c Institute for Bone and Joint Research, Kolling Institute, Sydney Medical School, University of Sydney, Australia b

A R T I C LE I N FO

A B S T R A C T

Keywords: Cadaveric Animal model Surrogate Finite element model Closed column Spinal cord injury

Background: The mechanical interactions occurring between the spinal column and spinal cord during an injury event are complex and variable, and likely have implications for the clinical presentation and prognosis of the individual. Methods: The engineering approaches that have been developed to better understand spinal column and cord interactions during an injury event are discussed. These include injury models utilising human and animal cadaveric specimens, in vivo anaesthetised animals, finite element models, inanimate physical systems and combinations thereof. Findings: The paper describes the development of these modelling approaches, discusses the advantages and disadvantages of the various models, and the major outcomes that have had implications for spinal cord injury research and clinical practice. Interpretation: The contribution of these four engineering approaches to understanding the interaction between the biomechanics and biology of spinal cord injury is substantial; they have improved our understanding of the factors contributing to the spinal column disruption, the degree of spinal cord deformation or motion, and the resultant neurological deficit and imaging features. Models of the injury event are challenging to produce, but technological advances are likely to improve these models and, consequently, our understanding of the mechanical context in which the biological injury occurs.

1. Introduction to injury biomechanics of spinal cord injury Excepting injuries caused by direct penetration (e.g. knife or bullet), traumatic spinal cord injuries (SCI) occur when the spinal canal is disrupted in a manner that allows excessive cord deformation or contact between the soft tissue or bony elements of the spinal column and the delicate spinal cord, overlying dura, and/or nerve roots. In adults, 55% of spinal injuries occur in the cervical spine; the remainder are equally distributed amongst the thoracic, thoracolumbar and lumbosacral regions (Sekhon and Fehlings, 2001). The most common mechanisms of spinal injury leading to SCI are fracture dislocations (40%) and burst fractures (30%), followed by less common fractures, dislocation and soft tissue injuries (Sekhon and Fehlings, 2001). Engineering approaches to understanding the mechanisms of the bone-cord interaction seek to provide quantitative understanding of the extrinsic and intrinsic factors contributing to the spinal column disruption, the degree of spinal cord deformation or motion, and the resultant neurological deficit and imaging features. Extrinsic factors are

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the external loads applied (including magnitude, rate, direction, and manner (inertial vs. contact)). Intrinsic factors include, for example, initial spinal posture, bone and soft tissue quality and degeneration, anatomical geometry (particularly with respect to canal and cord dimensions), and muscle activation. A number of modelling domains are used in the pursuit of this understanding; experimental modelling of injuries with cadaveric tissue (human or animal), modelling with live (anaesthetised) animals, finite element or other computational modelling, and inanimate physical models of sub-systems. Where other reviews may take a wider lens to a specific modelling domain (e.g. a broader review of in vivo animal models of SCI), this review focuses on the aspects of spinal column injury leading to SCI from a wide range of approaches. The aim is to provide an overview and commentary of the numerous engineering approaches that have been used to understand the relationship between extrinsic and intrinsic parameters of spinal column injury related to the cord injury. We include discussion of the challenges associated with using and interpreting these models, and the insights gained from these models.

Corresponding author at: MMBL, level 10, Kolling Building 6, RNS Hospital, St Leonards, NSW 2065, Australia. E-mail address: [email protected] (E.C. Clarke).

https://doi.org/10.1016/j.clinbiomech.2018.03.019 Received 25 August 2017; Received in revised form 16 February 2018; Accepted 24 March 2018 0268-0033/ © 2018 Elsevier Ltd. All rights reserved.

Please cite this article as: Jones, C.F., Clinical Biomechanics (2018), https://doi.org/10.1016/j.clinbiomech.2018.03.019

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compliant silicone elastomer tubing, filled with 0.9% saline solution to create a resistive element. An electric current was introduced into the saline-filled tubing at one end, and travelled towards a ground electrode at the opposite end. Seven electrode pairs, oriented across the tubing diameter, measured the resistance of the fluid (proportional to the area of fluid existing at that level at any given instant) at their respective locations. The electrodes are intended to be aligned with vertebral levels. The transducer was reliable between cross-sectional areas of 20–100 mm2 (with a nominal undeformed cross-sectional area of 125 mm2), with a resolution of ± 5 mm2. The SCOT was subsequently used in a series of studies (Table 1). Nuckley et al. (2007) instrumented baboon cervical segments with SCOT transducers to measure dynamic changes in the spinal canal when the specimens were dynamically compressed. The average peak occlusion (as a percentage of pre-injury canal diameter) was 52%, and was not dependent on ageor spinal level. In human thoracic spine specimens subjected to unconstrained and constrained axial compression, extension-compression and flexion-compression displacements, it was observed that constrained motion led to greater canal occlusion (due to more bony injury), while occlusion was lower for eccentric- than axialand flexion-compression loading (Zhu et al., 2008). In cervical segments loaded in dynamic eccentric axial compression, occlusion increased during loading and was highest at spinal levels where fractures or dislocations occurred (Van Toen et al., 2014). Peak occlusions were up to 76 and 33% for the low and high eccentricity groups respectively, and corresponded to hard and soft tissue injuries, respectively. Chang et al. (1994) studied dynamic changes in canal geometry during cervical spine injury using a fluid-filled tube spinal cord surrogate. Transient changes in the cross-sectional area of the spinal canal were related to the measured change in fluid pressure (Table 1). However, this method could not relate the canal occlusion to a particular spinal level because deformation of the tube caused the same pressure rise regardless of location. The same device was used in spine specimens loaded axially at low and high rates of equivalent energy (Tran et al., 1995). High-rate loading produced significantly higher canal pressures and therefore higher canal occlusions due to burst fracture (Table 1). Carter et al. (2000) used the pressure tube to study cervical spines subjected to axial loading in either a materials testing machine (“slow”) or drop tower (“fast”). They produced burst and anterior wedge compression fractures for the fast and slow methods, respectively. Transient peak canal occlusion was higher for burst fractures, than compression fractures (Table 1). Post-trauma occlusion was significantly lower than the transient for burst fractures, but not for compression fractures (for which some specimens registered no transient occlusion). Although the occlusion measure could not be definitely attributed to the level of injury, the clinical implication of this finding is that, for burst fractures, the bony fragments are maximally projected into the canal during the event, but this projection reduces after the insult, thus post-injury imaging of the bone alone may not reveal the full potential for SCI. Panjabi et al. (1995) describe a method of directly measuring linear mid-sagittal occlusion of the thoracolumbar canal at individual spinal levels (Table 1). The transducers consisted of a U-shaped strip of spring steel, the ends of which abutted the mid-sagittal anterior and posterior canal, and to which a uni-axial strain gauge was applied centrally. Bending of the steel strip, caused by shortening of canal depth, led to a change in strain reading from the gauge, and each transducer was calibrated against a gold standard displacement transducer. The transducers, placed at the superior endplates of T12 and L1, and mid-disc at T12-L1, were used in a dynamic model of burst fracture, showing that dynamic canal occlusion was ~85% more than post-injury static measurements of occlusion. The reported limitations of this approach were that the wires sometimes shorted due to fluids in the canal (although this presumably could be overcome by adequate coating of the strain gauge with impermeable silicon or polyurethane), and that the transducers could be dislodged immediately after peak occlusion due to bony fracture.

2. Experimental modelling of SCI within the spinal column Measuring dynamic spinal canal encroachment and/or spinal cord deformation during simulated, closed-column, spinal trauma can provide insight into the mechanisms and thresholds of SCI, however this is technically difficult. Several approaches have been taken to provide measures of “spinal cord” deformation within a cadaver spinal specimen during a simulated injury event. These can be broadly divided into five groups: measuring canal pressure, measuring spinal canal occlusion with custom transducers or via visual means, deriving canal occlusion from rigid body kinematics, visualising deformation of a physical surrogate, and measuring cord motion via markers. While the method of inducing spinal loading (e.g. materials testing machine, drop tower, acceleration sled, pneumatic impactor) and the specific specimen preparations to provide appropriate spinal posture and alignment (e.g. muscle replication, wedges to orient specimen etc.), simulation of other body parts (e.g. mass to simulate head mass and moment of inertia), are of great importance, they are outside the scope of this review to describe in detail. In this section we concentrate on the methods that have been developed to derive information about spinal cord and nerve root trauma from human and animal cadaver models of spine trauma. 2.1. Measuring spinal canal pressure Pressure transients within the canal, or within a surrogate cord placed in the canal, have been measured during experiments subjecting cadaver spinal segments to loads simulating external forces present during injury events. Pintar et al. (1995) subjected head-neck complexes with an instrumented gelatin surrogate cord to axial compressive loading with anterior, posterior or no eccentric alignment (Table 1). The cord filled the spinal canal, and the sensors were placed against the anterior canal wall. Compression and flexion-compression injury mechanisms led to burst and wedge compression fractures of the mid-cervical region, while extension compression mechanisms led to anterior ligament and disc injures. Local pressure change was generally proportional to injury severity (Pintar et al., 1995) (Table 1). Xie et al. (2001) removed the spinal cord (leaving dura intact) from thoracolumbar specimens, filled the dural cavity with saline and placed a pressure transducer in the fluid at the L1 level. The canal ends were sealed, with the specimen embedded such that L1 and its adjacent discs were exposed. Specimens were subjected to axial impacts to produce burst fractures (Table 1). Two pressure profiles were observed; the first exhibited positive pressure only, the peak ranging from 29 to 322 kPa, and corresponded generally with bone and disc fragments that contacted the dura. The second had both positive and negative pressure waves, the peak ranging from 65 to 81 kPa, and seemed to correspond to disc deformation without bony damage. While pressure transients that are transmitted to the cerebrospinal fluid or spinal cord, even without physical bony or soft tissue cord impingement, are a potential SCI mechanism, little is known about the magnitude, duration, or frequency, required of such transients to cause temporary or permanent tissue damage (Jones et al., 2013). Had the pressure changes in these models been calibrated to provide information on local surrogate cord deformation (see Section 2.2, SCOT sensor), it may have been possible to relate the bony and disc/ligament injuries observed to risk of SCI. 2.2. Measuring spinal canal occlusion with custom transducers or visually Raynak et al. (1998) first described their “neural space occlusion transducers, SCOT”, in 1998. These electromechanical sensors were designed to fully occupy the spinal canal and intervertebral foramen of cadaveric spines, replacing the spinal cord and nerve roots, to measure dynamic geometric changes. The sensors were constructed of highly 2

3

Mid-sagittal canal depth transducers Visualise bony occlusion

Panjabi et al. (1995) Wilcox et al. (2002) Panjabi et al. (2006)

Rigid body kinematics

Rigid body kinematics

Rigid body kinematics

Rigid body kinematics

Radiopaque surrogate spinal cord

Custom radiopaque markers

Tominaga et al. (2006)

Ivancic et al. (2007)

Ivancic et al. (2006)

Ito et al. (2004)

Saari et al. (2011)

Kroeker and Ching (2013)

Rigid body kinematics

Fluid-filled tube (pressure-to-CSA)

Carter et al. (2000)

High speed camera tracking sagittal plane vertebral body kinematics, geometric derivation of sagittal plane canal pinch diameter (CPD). Silicon elastomer (QMSkin 30, Quantum Silicones) cord dosed with barium sulfate, sagittal diameter visualised with high speed x-ray. Barium sulphate markers (approximately 3 mm diameter) in carrier paste injected into post-mortem spinal cord, visualised with fluoroscopic c-arm.

High speed camera tracking sagittal plane vertebral body kinematics, geometric derivation of sagittal plane canal pinch diameter (CPD).

High speed camera tracking sagittal plane vertebral body kinematics, geometric derivation of sagittal plane canal pinch diameter (CPD)

High speed camera tracking sagittal plane vertebral body kinematics, geometric derivation of foraminal dimensions

Strain gauge mounted on U-shaped spring steel. Maximum error 0.2 mm. Images of canal produced with mirrors and high speed camera. High speed camera tracking sagittal plane vertebral body kinematics, geometric derivation of foraminal dimensions

Canal cross-sectional area via change in tube fluid pressure; non-specific to spinal level.

Hyperextension injuries; peak cord compression:19–78%; peak compression length: 5.1–29.8 mm; 8–15 ms after initial impact. Complete dislocation occurred, most commonly at atlanto-occipital joint (mean load 1.95 kN). Mean inter-level cord-to-canal coupling ratio 0.54. Failure by complete dislocation usually at C0–1. Markers injected at each cervical level in 1-h postmortem NHP, tensile loading to column failure in a custom apparatus.

Rear impact with head turned causes foraminal narrowing sufficient for potential ganglion compression in patients with a nonstenotic foramen at C5–6 and C6–7. Dislocations in 10/12 specimens at 6.5–12.5 g. Peak CPD narrowing: C3/4 7.2 (SD3.0) mm; C5–6 6.4 (SD3.6) mm; C7-T1 5.1 (SD2.5) mm. Dynamic CPD < post-impact CPD (0.2 mm vs. 6.2 mm). Greatest CPD narrowing occurred at C0-dens (8.5 mm, or 25.9%, occurred at C0-dens at 10 g); generally increased with impact severity. No risk for cord compression in frontal impact to 10 g with or without canal stenosis. Narrowest CPD at C5/6 at 6.5 g: 3.5 mm < neutral posture CPD. SCI during whiplash unlikely with normal canal diameters.

Rear impact could cause foraminal width and area narrowing to compress nerve roots and ganglia.

Post-injury occlusion < peak occlusion during impact event

High rate loading produced burst fractures and higher canal pressures and occlusions; low rate loading produced compressive fracture without significant canal occlusion. Peak canal occlusion: 65–75% for burst fractures; 0–32% for compression fractures. Post-trauma occlusion < peak transient for burst fractures. Dynamic canal occlusion ~85% > post-injury static occlusion

Observed transient changes in canal cross-sectional area during injury event.

Peak occlusion up to 76% (low eccentricity) with bony injury and 33% (high eccentricity) with soft tissue injury.

Positive pressure only, peak 29–322 kPa: bone/disc fragment contact with dura; negative/positive pressures, peak 65–81 kPa: disc deformation without bony damage The average peak occlusion (% pre-injury canal diameter) was 52; was not dependent on age (1–30 human-equivalent years old) or spinal level. Constrained motion results in greater canal occlusion due to more bony injury.

Lordosed cervical spine, head-first dynamic impact

Whole cervical spines, incremental inertial loading (rear impact) with sled, (3.5, 5, 6.5 and 8 g).

Whole cervical spines, incremental inertial loading (frontal) with sled (4, 6, 8, and 10 g).

Thoracolumbar burst fracture, transducers placed at T12 and L1 superior endplates and mid disc. Experimental burst fracture; bovine thoracolumbar segments Cervical spine (Occ-T1), incremental inertial loading with bench-top sled apparatus (baseline 2 g, then incremental 3.5, 5, 6.5, and 8 g until failure observed). Cervical spine (Occ-T1) with axial rotation of head, incremental inertial loading with sled (baseline 2 g, then incremental 3.5, 5, 6.5, and 8 g until failure observed). Cervical FSUs (C3/4, C5/6, C7/T1), incremental inertial loading (frontal impact configuration) with sled until bilateral dislocation

Axial compression of non-lordosed cervical spines at high (drop tower) or low (MTM) rates.

Spinal canal crosssectional-area (fluidfilled tube) Fluid-filled tube (pressure-to-CSA)

Chang et al. (1994)

[SCOT]

Canal cross-sectional area via change in tube fluid pressure; non-specific to spinal level.

Spinal canal occlusion

Van Toen et al. (2014)

Thoracolumbar burst fracture (three-vertebrae bovine), low and high rate axial loading of equivalent energy

Spinal canal occlusion

Zhu et al. (2008)

Thoracolumbar specimens, axial impact (5.4 m/s, 20 ms duration, loads not reported) to produce burst fractures

Canal cross-sectional area via change in tube fluid pressure; non-specific to spinal level.

Spinal canal occlusion

Nuckley et al. (2007)

Tran et al. (1995)

Spinal canal pressure

Xie et al. (2001)

More severe injuries led to higher local pressure changes: unaffected levels 0–300 kPa; minor injuries 350–750 kPa; significant fractures or dislocation 500–2000 kPa.

Head-neck complexes; axial compression loading with anterior, posterior or no eccentric alignment, in a materials testing machine at 300 or 600 cm/s.

Three-vertebra cervical segments (Oc–C2, C3–C5, and C6–T1) post-mortem baboon, dynamic axial compression to 70% strain. Thoracic segments, loaded in unconstrained and constrained axial compression, extension-compression and flexion-compression. Three-vertebra cervical segments loaded in dynamic (0.5 m/s) eccentric axial compression (1 or 150% of lateral vertebral body diameter). Thoracolumbar burst fracture

Collagen encased gelatin cord instrumented with 7 piezoelectric pressure sensors (1 mm thick, 10 mm2, approximately 15 mm apart). Gelatin mixture varied to provide response similar to in vivo feline spinal cords (with dura and CSF) subjected to weight drop. Intact dura filled with saline, pressure transducer at L1.

Spinal canal pressure

Pintar et al. (1995), Pintar et al. (1996)

Major findings

Injury model

[SCOT] Saline-filled silicon elastomer tubing, electrode pairs aligned with vertebral levels measured change in resistance corresponding to level specific cross-sectional area. [SCOT]

Methodology

Category

Author/year

Table 1 Experimental modelling of SCI within the spinal column. *Human unless otherwise noted. Acronyms: CSF: cerebrospinal fluid; MTM: materials testing machine; FSU: functional spinal unit; CPD: canal pinch diameter; NHP: non-human primate.

C.F. Jones, E.C. Clarke

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area of the foramen). Geometric relationships, assuming each vertebrae was a rigid body, between the body markers and foramen points were defined with the lateral radiograph, and the translation of foramen points was defined throughout the motion, according to these relationships, using kinematic calculations. The study showed that rear impact could cause foraminal width and area narrowing sufficient to compress the cervical nerve roots and ganglia, especially in those with foraminal spondylosis. Similar techniques were used to investigate foraminal narrowing during rear impacts when the head is turned (Tominaga et al., 2006) (Table 1). Ivancic et al. (2007) estimated dynamic canal pinch diameter (CPD) narrowing in functional spinal units subjected to incrementally increasing accelerations until bilateral facet dislocation (BFD) occurred (Table 1). Rigid body transformation of kinematic data, with lateral xrays defining the initial anterior and posterior limits of the canal relative to the vertebral body markers, was used to compute CPD in the neutral posture, during dislocation, and post-impact. CPD was defined as the distance between the inferior corner of the spinolaminar line of the upper vertebra and the posterosuperior corner of the lower vertebral body. Dislocation occurred in 10/12 specimens with peak CPD narrowing of 7.2, 6.4 and 5.1 mm, at the C3–4, C5–6 and C7-T1 levels, respectively. Dynamic CPD narrowing was significantly higher than post-impact CPD narrowing, suggesting that spinal canal narrowing observed on post-trauma medical imaging of BFD patients is likely significantly less than the dynamic narrowing that occurred during the dislocation. Using a similar approach to estimate CPD, but with whole cervical spine specimens subjected to (inertial) frontal impact loading, Ivancic et al. (2006) established the chronology of CPD at each level, and the relative inter-level CPD magnitudes for this injury mechanism. Any facet dislocations or soft tissue damage that occurred were not reported. In general, the peak CPD narrowing at C0-Dens and C1–C2 occurred immediately following the peak T1 acceleration during intervertebral extension of the upper cervical spine, while the peak CPD narrowing at C2–C3 to C6–C7 occurred later during peak intervertebral flexion. This agrees with the two-phase protraction-flexion kinematic pattern noted for the cervical spine in frontal impacts (Deng et al., 1998; Thunnissen et al., 1995). CPD narrowing was generally proportional to impact severity (Table 1). With some corrections for the estimated soft tissue contribution to canal diameter, the authors concluded that there is no risk for cord compression during frontal impact (of up to 10 g acceleration) in individuals with normal canal diameters or in those with cervical canal stenosis. The same model and procedure was used to determine CPD in rearimpact “whiplash” accelerations (Ito et al., 2004). It was hypothesised that the hyperextension observed in the lower cervical spine in the initial phase of whiplash following rear impact would cause the spinal canal diameter to decrease and put the cord at risk of injury. The narrowest CPD was observed at C5/6 and was 3.5 mm narrower than the neutral posture CPD. They concluded that SCI during whiplash is unlikely in individuals with average normal canal diameters.

Wilcox et al. (2002) devised a method of visualising bony occlusion during simulated burst fractures. They used a series of mirrors mounted above and below the spinal canal (spinal cord, dura and fat cleared from the canal), in combination with a light source and high speed video camera, to create binary images showing the bony fragment being retropulsed into the canal. From this, the change in spinal canal crosssectional area was derived. This is the only reported method that directly visualises the column and does not introduce measurement devices or surrogate spinal cords into the canal, which may affect bone fragment motion. However, it is limited to tests with a small number of spinal segments, and those in which vertical canal alignment is maintained to ensure the passage of light. Importantly, the study provided measures of bone fragment velocity, mass and cross-sectional area (which informed the design of in vivo animal (Jones et al., 2012b) and physical (Jones et al., 2012a) models of SCI) and provided evidence that in clinical burst fracture the post-injury occlusion is likely lower than the peak occlusion (Wilcox et al., 2002).The results from this approach were later combined with a model in which the canal was filled with gelatin and an embedded pressure transducer, to determine transient pressures, and the results used to develop and validate a finite element model (see Section 4) (Wilcox et al., 2003; Wilcox et al., 2004). With the latter model, fragment propagation and recoil were shown to be dependent on fragment velocity and the presence of spinal cord and dura (Wilcox et al., 2003; Wilcox et al., 2004). The calibration of the fluid-filled transducers of Raynak et al. (1998) and Chang et al. (1994), and the level specific strain gauge-based sensors of Panjabi et al. (1995), lead to the first measures of transient canal dimensions, which (in the case of burst fractures) provided evidence that SCI severity may not be reliably derived from post-injury bony imaging. This was corroborated by the “direct visualisation” technique of Wilcox et al. (2002). That few of these techniques have been used over a large number of studies may point to practical difficulties in their implementation or their suitability for use in combination with other experimental techniques. 2.3. Deriving canal occlusion from experimental kinematics A number of authors have derived canal occlusion or intervertebral foramen occlusion using optical marker tracking and rigid body kinematics. These techniques are not usually suitable for experiments in which fracture is expected, because at the time of fracture the elements being tracked may cease to be rigid bodies (and therefore the marker motion may not faithfully represent the element motion). Thus, for the spine, this approach is limited to injury mechanisms causing soft tissue disruption, dislocation of the facets, or disruption of the intervertebral disc, in which case bony elements are not deformed or fractured. Although there is currently little published data for the cervical spine, up to 14 degrees of sagittal rotation of the facet tips relative to the vertebral body have been observed in lumbar spine specimens during experimental flexion (Green et al., 1994), thus caution should be applied to rigid body assumptions where the posterior elements are concerned. The kinematic approach may be used with “digitised” points to track virtual anatomical markers, or the kinematics may be applied to a 3D model derived from CT, so that the occlusion can be visualised, and anatomical landmarks not able to be physically digitised may be tracked. A further limitation of the rigid body kinematics approach is that soft tissue bulging into the canal dynamically (e.g. posterior longitudinal ligament and ligamentum flavum) cannot be accounted for and these may increase the severity of cord compression and neurologic injury. Using high speed video, Panjabi et al. (2006) tracked the motion of markers attached to the vertebral bodies of cervical spines during a dynamic rear-impact simulation (Table 1). Using specimen-specific radiographs and literature-derived quantitative anatomy of the cervical foramen, six foraminal points were mathematically reconstructed for each functional spinal unit (and used to define the width, height and

2.4. Visualising deformation of a physical surrogate spinal cord Saari et al. (2011) used a radiopaque surrogate spinal cord developed and validated by Kroeker et al. (2009) to visualise sagittal spinal cord deformation within the lordosed cervical canal in head-first dynamic impact simulations. The re-usable radiopaque cord was moulded from a silicon elastomer, and validated with mechanical testing in quasistatic tension and compression (Kroeker et al., 2009), and dynamic compression (Jones et al., 2008). The radiopaque cord was visualised with high speed x-ray, and the resulting images manually analysed to determine dynamic sagittal cord diameter. Transient cord compressions of 19–78% and compression spans (along the cord) of 5.1–29.8 mm, occurred 8–15 ms after initial impact. Larger compressions were associated with atlantoaxial dislocation and smaller 4

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available for cord” is clinically considered to influence risk of SCI, and likely influences the bone-cord interaction. In addition to providing clinical understanding of the utility of postinjury imaging for prediction of neurological outcome, the results of the models described above have been used to validate finite element (FE) models, and as a basis for the design of in vivo animal models of SCI, thus providing important information for the advancement of basic and applied research in spine injury and SCI.

compressions associated with hyperextension injuries (via a pincer mechanism), and no specimens had residual compression of the cord when the spines were returned to a neutral posture after injury. This model did not simulate the dura and CSF, and the cord was not tethered in the canal (Saari et al., 2011); the latter may lead to substantial cord mobility and as a result any increase in canal length may not result in cord lengthening and narrowing that would occur in vivo. 2.5. Alteration of the natural spinal cord with markers

3. In vivo modelling with live anaesthetised animals Kroeker and Ching (2013) injected custom barium sulphate radiopaque markers at each cervical level into the spinal cords of postmortem non-human primates (NHP) (Table 1). The cervical column was subjected to tensile loading until failure, and the displacement of the spinal cord (relative to lead beads adhered to the laminae at each level), was tracked with a fluoroscope. Failure resulted in complete dislocation of vertebrae, most commonly at the atlanto-occipital joint. Level-specific cord-to-canal coupling ratios, with a mean ratio of 0.54, indicated that under tension-extension loading, the adult NHP spinal cord strains less than the column. It should be noted that this technique relied on having very fresh spinal cord; it is unlikely to produce accurate results for fresh-frozen or cadaveric tissue obtained later after death (such as that used in human cadaveric models of injury), because the postmortem spinal cord degrades quickly and loses its mechanical integrity.

In vivo animal models of traumatic SCI are used to determine functional and pathological SCI outcomes following controlled injury simulations in living tissues. These models can be classified as open- or closed-column. The term “open-column” generally refers to models where a portion of the spinal column is removed to expose the spinal cord that is then directly injured (e.g. compressed, contused/impacted or cut). The term “closed-column” generally refers to models that perform loading on the intact spinal column (e.g. dislocation, distraction), which indirectly injures the spinal cord within. However, there are some exceptions; for example, some distraction models include an osteotomy to expose the spinal cord but apply the loading to the spine (e.g. Owen et al., 1988), and the SC-STRAPPER model directly injures the spinal cord via the intervertebral foramen without opening the spinal column (da Costa et al., 2008). For this review, any model that includes an osteotomy is classified as open-column because spine/ spinal cord coupling could be compromised. Also, although the SC-STRAPPER model leaves the spinal column intact, for the purpose of this review, we classify this model as an open-column model because the loading is applied directly to the spinal cord itself. We focus on closed-column models in this review because they are unique in enabling investigation, in living tissues, of mechanisms of spinal injury leading to SCI. Generally, closed-column models apply various forms of dislocation or distraction to the spinal column; however, it is worth noting the unique canine transverse impact model reported by Wennerstrand et al. (1978). Their transverse impact model was rather similar to contemporary spinal cord impactor (contusion) models, but the impact was applied to the spine rather than to the spinal cord. Although most of the animals (4/6) reportedly had no vertebral column “injury” (e.g. fracture), all but one had histological damage to the spinal cord at 3 h post-injury. They used a small number of animals and varied the impact displacement and speed simultaneously, making it difficult to determine which aspects of the column injury produced cord injury, and it appears there are no follow-up studies using the same model to further investigate this mechanism of SCI.

2.6. Challenges There are advantages and disadvantages associated with each of the described approaches to quantifying spinal cord involvement during simulated spinal trauma. The ultimate cadaveric model of spinal column and cord injury would provide a continuous and three-dimensional measure of spinal cord dimensions throughout the simulated trauma. The various sensors developed to reside within the canal produce either non-localised or discrete measures of canal occlusion, while the radiopaque surrogate cord approach provides continuous measure of cord size along its length, but only in one plane. Measures of canal pressure provide an indication of the dynamic stress imparted on a surrogate cord, but not its physical dimensions throughout the event. These techniques have provided valuable insight into the relative value of post-burst fracture bony imaging (x-ray and CT) but while measures of canal occlusion (which may be converted to cord compression) may be related to neurological deficit via existing animal model data, measures of canal pressure alone are difficult to relate to cord injury risk until appropriate relationships are derived in animal models. Rigid body kinematics assumptions may be violated upon injury if the rigid body of interest is fractured (or before injury if bony deformations are significant), and thus inferring canal or foramen dimensions using this method is largely limited to inertial loading mechanisms. Markers must be light-weight (to minimise effect on body inertia) and fixed to the bone in a manner that does not induce damage and pre-dispose the specimen to non-physiologic injury patterns, yet be able to withstand high accelerations without becoming detached. Any technique that relies on visual tracking requires relatively expensive high speed cameras or high speed x-ray with adequate resolution. Techniques that visualise the canal directly from above or below, are only suited to specimens that remain axially aligned throughout the test. Any surrogate cord developed must have similar dynamic mechanical properties to the native cord (and dura) in order that the interaction between the bony elements and the surrogate material are biofidelic. This is challenging given the paucity of data available for the fresh human spinal cord tested at high speeds. In addition, no cadaveric models have yet attempted to include a dura or CSF simulant in the system, which have been shown to influence cord compression in a number of physical models (see Section 5). The “fit” of the surrogate cord within the canal is another important consideration, as the “space

3.1. Distraction The most common closed-column model of SCI is the distraction mechanism, in which the spine and contained spinal cord are loaded in axial tension. Spinal distraction is clinically relevant to spinal surgery with traction (e.g. for spinal deformity correction (Lewis et al., 2014)), and tension-extension injuries of the spine (which can occur when airbags are deployed under the chin (Eppinger et al., 1999) and in neck hyperextension from rear impact (Lang, 1971; Panjabi and Myers, 1995)). Spinal distraction studies have provided insight to the coupling between the spine and spinal cord, by using radio-opaque markers and fluoroscopy/radiography. Kroeker and Ching (2013) performed spinal distraction in primates (this is described in detail in Section 2.5 because they used cadaveric tissue). Maiman et al. (1989b) performed quasistatic distraction in the cat cervical and lumbar spine below the threshold expected to produce spine/SCI (Table 2). Despite using different species and loading rates, both studies reported mean spinal cord/spine coupling ratios of 53–54%. This insight is useful for predicting spinal cord strains based on vertebral column strains. 5

Category

Spinal distraction

Spinal distraction

Spinal distraction

Spinal distraction

Spinal distraction

Distraction, dislocation and contusion

Distraction, dislocation and contusion

Distraction, dislocation and contusion

Dislocation

Dislocation

Dislocation

Dislocation

Dislocation (and contusion)

Author/year

Maiman et al. (1989a)

Maiman et al. (1989b)

Dabney et al. (2004)

Hong et al. (2016)

Wu et al. (2017)

Choo et al. (2007)

Choo et al. (2008)

Chen et al. (2016)

Fiford et al. (2004)

Clarke and Bilston (2008)

Clarke et al. (2008)

Lau et al. (2013)

Bhatnagar et al. (2016b)

Major findings

Mean coupling ratio cord-to-spine 53% for low/high loads (78% for intermediate loads); inconsistent across spinal levels; useful for predicting cord strain from vertebral strain Quasi-static distraction applied until 50% or 95% Feline; rigid hooks at iliac crest and distraction applied Reduction in evoked potential correlated with distraction force reduction in evoked potential amplitude; (coupling location not specified but presumably the head). applied; tissue damage correlated with ultimate neurologic status Motorised distraction device; length, speed or duration Rodent; vertebral distraction applied via sublaminar hooks at Distraction displacement affected SCI but not speed and short (seconds) of distraction altered T9 and T11 duration of distraction; Porcine; vertebral distraction applied via pedicle screws at T13 Spinal distraction of > 70% of segmental vertebral height may Distraction-displacement target held (for minutes); and L1 cause irreversible SCI; longer duration distraction monitored for loss and recovery of motor evoked (mean > 10 min) may worsen SCI potentials. Varying degrees of spinal distraction applied (10%, 20% Rabbit; vertebral distraction applied via clamps secured around Spinal distraction leads to spinal cord ischemia, which is a primary pathological mechanism of distraction SCI or 30% of L1-L3 distance); pathological mechanisms of T12 and L4 vertebrae SCI investigated Rodent; vertebral distraction or anterior dislocation applied via Different mechanism of spinal injury cause different patterns of Multi-mechanism injury model applied high speed primary (immediate) SCI histopathology; in general contusion and clamps secured around C3/4 and C5/6; or contusion applied flexion-distraction (mean 91.9 cm/s), fracturevia impactor directly on spinal cord while C4 and C5 vertebrae dislocation more focal and more severe than distraction dislocation (95.1 cm/s) or cord contusion (96.7 cm/s) rigidly clamped as one unit Rodent; vertebral distraction or anterior dislocation applied via Different mechanisms of spinal injury cause different patterns of Multi-mechanism injury model applied high speed secondary (3 h post-injury) SCI histopathology; in general contusion clamps secured around C4 and C5; or contusion applied via flexion-distraction (mean 111.2 cm/s), fractureand dislocation more focal and more severe than distraction impactor directly on spinal cord while C4 and C5 vertebrae dislocation (95.6 cm/s) or cord contusion (99.8 cm/s) rigidly clamped as one unit Rodent; vertebral distraction or anterior dislocation applied via Different mechanism of spinal injury cause different patterns of Multi-mechanism injury model applied high speed functional deficit/recovery; in general dislocation caused greater clamps secured around C4/5 and C6/7; or contusion applied flexion-distraction (mean 129.8 cm/s), fracturedislocation (87.0 cm/s) or cord contusion (120.6 cm/s) via impactor directly on spinal cord while C5 and C6 vertebrae functional deficits than distraction and contusion (despite the lower speed). rigidly clamped as one unit Rodent thoracolumbar fracture-dislocation at varying Graded SCI severity produced by increasing speed and displacement Rodent; lateral vertebral dislocation applied via clamps speeds (57–127 mm/s) and displacements (3.2–7.5 mm/ secured around L2 and T13 vertebrae (T12 was specified but of dislocation (varied simultaneously); spinal cord regions that were predicted to have greater longitudinal deformation during spinal T13 was used) s) injury had greater axonal injury Rodent; lateral vertebral dislocation applied via clamps Adults had fracture-dislocations, neonates had no fractures; despite Rodent thoracolumbar dislocation comparing adult scaled spinal column injury severities in adults and neonates, SCI (259 mm/s) and neonate (93.4 mm/s, scaled by vertebral secured around T12 and L1 vertebrae severity was generally worse in neonates when normalised to body height) account for their smaller size Rodent; lateral vertebral dislocation applied via clamps Vertebral fracture occurred at a lower displacement in lateral Comparison of lateral and anterior fracture-dislocation in rodents at 220 mm/s secured around T12 and L1 vertebrae dislocation; different directions of dislocation produced different spatial distributions of SCI pathology; SCI severity more severe in anterior dislocation Independent variation of speed and displacement of Rodent; lateral vertebral dislocation applied via clamps Age and displacement (but not speed) of dislocation are key factors dislocation; comparing adults and neonates secured around T12 and L1 vertebrae in the severity of acute SCI High-speed dislocation or contusion in MRI scanner to Rodent; anterior dislocation applied via clamps secured around Anterior dislocation caused a dorso-ventral tensile strain “band” measure cord strain during injury C4/5 and C6/7 that extended laterally across the central region of the cord

Experimental model

Quasi-static distraction; displacement in spine and spinal Feline; rigid hooks at iliac crest and distraction applied cord tracked to measure coupling (coupling location not specified but presumably the head).

Methodology

Table 2 In vivo modelling of closed column spinal injury with live anaesthetised animals.

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(Choo et al., 2009) and the primary and secondary spinal cord pathology were initially provided by Choo et al. (2007, 2008) (Table 2). Since the dislocation mechanism produced an unstable spinal fracture, the UBC group designed and biomechanically tested a rodent spinal fixation system (Shahrokni et al., 2012) and were then able to perform survival studies to compare behavioural outcomes for the three injury mechanisms (Chen et al., 2016). Together, these studies have demonstrated that different spinal column injuries lead to very different SCI histopathology and behavioural outcomes. This is an exceptionally important finding for the field because: (1) it suggests that SCI therapies may need to be customised for different injury mechanisms; and, (2) it demonstrates that the contusion model, which is one of the most common models used for experimental research for SCI, produces cord pathology and functional deficit patterns that are different from the most common mechanism of SCI in humans. This may reduce the translatability of therapies developed or tested using these models. One of the limitations of dislocation models described above was that the motion of the cord was neither directly controlled nor visible during the injury event. To address this, the UBC group have produced a magnetic resonance imaging (MRI) compatible system for dislocation (and contusion) (Bhatnagar et al., 2014) so that imaging can be performed during the injury. In dislocation, the most obvious cord deformation “pattern” was a dorso-ventral tensile strain “band” that extended laterally across the central region of the cord (Bhatnagar et al., 2016b), which is somewhat different from the pattern of deformation predicted by Fiford et al. in lateral dislocation. In the future this MRI compatible model could be valuable in understanding the relationship between cord deformation and injury pathology during clinically relevant injuries. The Sydney and UBC models have been directly compared in the thoracolumbar spine; demonstrating that dislocation direction affected vertebral injury threshold (fracture occurring at a lower displacement in lateral than anterior), spatial distribution of cord pathology, and SCI severity (anterior more severe than lateral) (Clarke et al., 2008). This demonstrates that even with the same spinal column injury mechanism, simply changing the direction can significantly alter the SCI outcomes. Since fracture dislocation is the most common spinal injury mechanism in humans, and in vivo models using live anaesthetised animals have demonstrated that different injury mechanisms produce different functional and pathology outcomes, it is very important to refine and use these models to expand the range of options for experimental research and preclinical testing for SCI therapies.

Distraction models using specific displacement targets have demonstrated that SCI severity (across several outcome measures) is proportional to displacement (Dabney et al., 2004; Seifert et al., 2011; Wu et al., 2016; Wu et al., 2017) but not speed (Dabney et al., 2004) (Table 2). Other studies have applied continuous distraction until a specific injury severity threshold is achieved, e.g. reduced evoked potentials or structural (histological) and/or behavioural changes (Dabney et al., 2004; Hong et al., 2016; Maiman et al., 1989a). These studies have assisted in defining thresholds for various pathological processes during injury (e.g. damage to spinal cord microvasculature occurs before neurologic dysfunction (Wu et al., 2017)), and for demonstrating that evoked potentials (which could be measured in vivo clinically or experimentally) can be used to predict, monitor for, or specifically target, structural and/or functional SCI severity (e.g. Maiman et al., 1989a). All of these studies have been performed at relatively low or moderate speeds (order of 1–10 mm/s), whereas studies using another model (Chen et al., 2016; Choo et al., 2007; Choo et al., 2008) have comprehensively characterised the histopathology and functional outcomes of high speed (> 900 mm/s) spinal distraction injuries. Since the species and specific outcome measures used by various laboratories are different, it is difficult to compare between them – e.g. to compare distraction in one study performed at 1 mm/s with 900 mm/s in another study. 3.2. Fracture-dislocation In humans, spinal fracture-dislocations are the most common mechanism of traumatic SCI (Sekhon and Fehlings, 2001), and within this category, anterior dislocations (wherein the superior vertebra translates anteriorly) are most common, followed by lateral and then posterior dislocations (Tator, 1983). Animal models of spinal dislocation are therefore important to improve understanding of this most common mechanism, and for pre-clinical studies of potential therapies. To our knowledge, only two animal models of vertebral dislocation have been developed; the Sydney lateral dislocation model (Clarke and Bilston, 2008; Fiford et al., 2004; Lau et al., 2013), and the UBC multi-mechanism system, which can produce anterior dislocation (Chen et al., 2016; Choo et al., 2007; Choo et al., 2008; Choo et al., 2009). Both dislocation models have similar clamping and actuation methods, similar surgical approaches and similar patterns of SCI pathology; however, they have been utilised at different spinal levels and to address different specific research questions. Both models require resection of the facet joints to prevent perched facets upon realignment and both produce an endplate fracture. The first in vivo dislocation model of SCI was developed in rats by Fiford et al. (2004) (Table 2). They demonstrated that areas of the spinal cord experiencing greater longitudinal deformation exhibited axonal injury. They also demonstrated that graded severity of spinal cord hemorrhage and axonal injury could be achieved by altering the speed and displacement of the spinal dislocation (however, the speed and displacement were altered simultaneously). A later study using the same model altered speed and displacement independently, demonstrating that histological SCI severity was dependent on the displacement but not speed of dislocation (Lau et al., 2013), consistent with previous findings for spinal distraction (Dabney et al., 2004). The Sydney model has also been used to contrast spine/spinal cord injuries in adult and infant rats; where anatomically scaled dislocation displacement produced fractures in adults but not in infants and, when normalising for the different spinal cord sizes, the severity of histological cord injury was more severe in infants than adults (Clarke et al., 2009; Lau et al., 2013). This has provided insight to different patterns of injury that may occur in human adults and children during dislocation injuries. The UBC dislocation model is part of a multi-mechanism system that was developed to compare SCI produced by different spinal column injury mechanisms at high speeds. A detailed description of the system

3.3. Challenges Simultaneously, one of the greatest advantages and disadvantages of closed-column models is that the loading is applied to the spinal column, not the spinal cord. This is advantageous because it is more clinically relevant of human trauma and it allows investigation of the interaction between the spine and spinal cord during injury. However, it is also a disadvantage because the loading on the spinal cord is not controlled, and generally not measured or even visible, so injury severity can be variable (e.g. see the generally higher variability for dislocation than contusion in figures in Chen et al. (2016)). Some studies have attempted to visualise or measure cord motion during closedcolumn injuries; e.g. Fiford et al. (2004) performed a laminectomy in a small number of rats to view the spinal cord during lateral dislocation, and Kroeker and Ching (2013) used fiducial markers and fluoroscopy to measure spinal cord displacement during distraction. However, these required removal of spinal column tissue to access the spinal cord for visualisation and thus may have altered the interaction between the spine and spinal cord. Bhatnagar et al. (2016a) performed MRI scanning during anterior vertebral dislocation (held stationary in the displaced position) to measure internal spinal cord deformation in the injured position. This was an advancement in the field because this was the first visualisation and measurement of spine and spinal cord motion and 7

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(2002) developed a model of the human cervical spine and spinal cord to investigate spinal cord stresses and strains in flexion and extension injuries. They proposed that in flexion there may be higher stress in the anterior spinal cord than in the posterior. Wilcox et al. (2004) developed a FE model to investigate the burst fracture injury mechanism and validated it with experimental testing on specimens with the same geometry (see Section 2.2). In this study the spinal cord was included to allow for rebound of the bony fragments from the spinal injury, rather than to investigate the effects of the vertebral injury on the stresses and strains in the cord itself. This study provides insight to the mechanisms for a common spinal injury and guidance for the loading conditions in future studies of cord injury using the burst fracture mechanism. FE models with more complex anatomical structures have been developed to compare spinal cord stresses and strains from different spinal injury mechanisms. Greaves et al. (2008) developed a finite element model of the human cervical spine with spinal cord to compare cord strains arising from three different injury mechanisms: dislocation, contusion and distraction. Khuyagbaatar et al. (2016) also developed a FE model of the human cervical spine and spinal cord (albeit with material properties for the spinal cord and dura sourced from bovine tissues at low strain rate), comparing the same injury mechanisms as Greaves et al. Both studies reported a focal region of peak strain for contusion and dislocation, and diffuse peak strain for distraction. Khuyagbaatar et al. also observed that the region of high stress and strain was localised anterolaterally in dislocation but posterolaterally in contusion. They also observed that with increasing magnitude of dislocation, contusion or distraction, the effects on spinal cord stress, strain, and cross sectional area were more pronounced for dislocation, followed by contusion, followed by distraction. Further, when considering all three mechanisms, the reduction in spinal cord cross sectional area was highly correlated with peak stresses and strains, which implies that peak spinal cord stresses and strains could potentially be predicted from reduction in spinal cord cross sectional area. Finally, Russell et al. (2012) developed a complex FE model of the rat cervical spine and spinal cord to compare dislocation and contusion injures. Considerable effort was made to select and tune the material properties for the model, and to validate the model using forces recorded during in vivo experiments of the same injury mechanisms. Together, these FE models have provided engineering insight to the different structural and functional injuries observed in vivo for different spinal injury mechanisms. An exquisitely detailed head and neck FE model was developed by Kimpara et al. (2006) to investigate the relationship between restraint conditions in a vehicle and brain/spinal cord injuries. The authors report detail in selecting material properties, validating the model, and using the model to simulate real world head impact scenarios to refine injury criteria. While not specifically developed to be a model of spinal column injury, it provides insight to the conditions of head impact that could lead to spine or SCI, which provides useful insight for injury prevention strategies.

deformation during a controlled, closed column, SCI in living tissue. With further refinement this model will be a very useful tool in advancing our understanding of SCI mechanics, and will broaden the scope of data collection for preclinical testing of SCI therapies. Another challenge for these models is avoiding damage to the vertebrae via the clamps or relative motion between the vertebrae and clamps/hooks/screws during injury; model developers design very specific coupling methods to prevent this (e.g. Choo et al., 2009). The surgery to achieve secure coupling can also be extensive and invasive; some models require release of the paraspinal muscles to clamp the transverse processes (e.g. Choo et al., 2009; Clarke et al., 2008; Seifert et al., 2011). Some models require additional surgery, such as facet removal and transection of supraspinous/interspinous ligaments, to control the injury location and/or severity and to avoid post-injury vertebral misalignment (e.g. Choo et al., 2009; Clarke et al., 2008; Seifert et al., 2011), resulting in an unstable spine that requires fixation. The combination of these challenges means that these models are not widely used in comparison to open-column models; indeed only one study has (recently) maintained animals beyond a few hours post dislocation injury and obtained behavioural outcomes (Chen et al., 2016). Nevertheless, closed-column models are important for understanding injury mechanisms and thresholds, and testing therapies in clinically relevant injury modes during the final stages of pre-clinical testing. 4. Computational models of spinal column injury leading to SCI Finite element (FE) models provide another method for investigating the mechanisms of spinal injury leading to SCI. Such models allow impact conditions, geometry or tissue properties to be adjusted to investigate their potential effects on SCI outcomes, without the cost, time, ethical and logistical issues of animal studies. Several FE models have been developed, modelling a range of species, spinal levels and injury mechanisms, and with varying degrees of complexity (Table 3). For example, some models include a spine but no spinal cord (e.g. Imajo et al., 2009; Qiu et al., 2006), and others include a spinal cord but no spine (e.g. Kato et al., 2009; Li and Dai, 2009; Maikos et al., 2008). Slightly more complex models involve a spinal cord with a boundary to approximate the spinal canal (Czyz et al., 2008; Sparrey et al., 2016), and the most complex models include full spine and spinal cord anatomy, including dura and dural attachments (Sparrey et al., 2016). In this review, we focus on the mechanisms of spinal injury leading to SCI, so we discuss FE models that include both the spine and spinal cord, or models that perform spinal cord loading to mimic a clinical (or in vivo experimental) spinal injury. Some studies have used FE models of the spinal cord without complex spinal anatomy to investigate spinal column/cord injury. Although these models do not include spinal anatomy, they attempt to simulate real human clinical, or in vivo animal model, injuries, so they have very high-level validation and utility. Czyz et al. (2008, 2012)developed a FE model of the human cervical spinal cord to simulate subject-specific cases of SCI for patients with and without neurological deficit (Table 3). Their models predicted significant differences in stresses and strains between patients with and without deficit, and the spinal cord stresses and strains were also correlated with neurological deficit in the matched clinical cases of SCI. This demonstrates that FE models could potentially be used clinically to predict neurological deficits after injury, and may provide insight to injury conditions leading to neurological deficit. Sparrey et al. (2016) developed a non-human primate FE model of unilateral spinal cord contusion and compared it's outputs with in vivo experimental data for the same injury model, in order to provide insight for the development and refinement of the in vivo primate model while minimising animal use. For example, it was used to determine the effect of lateral alignment of the impactor, relative to the cord, on translation and rotation of the cord during injury. Other studies have developed FE models of the spine and spinal cord to investigate the mechanisms of specific spinal injuries. Scifert et al.

4.1. Challenges One of the main challenges in developing biofidelic FE models of injury is selecting appropriate material properties. Spinal cord injuries are generally high-speed events, and the properties of biological tissues are strain-rate dependent, so it is important to use material properties measured at appropriate strain rates when simulating injuries. Accurately measuring tissue properties at high speeds is technically challenging, so material properties are not always available for the specific tissues and conditions required for models of injury. Also, if several sources exist, material properties for the same tissue or structure can be quite variable, even when tested under similar conditions (Cheng et al., 2008). For these reasons, material properties may need to be specifically measured, or estimated or averaged from the literature, or a range of properties should be tested in a sensitivity analysis so that 8

Cord without complex spinal anatomy

Cord without complex spinal anatomy Simple spine and spinal cord

Simple spine and spinal cord

Czyz et al. (2012)

Sparrey et al. (2016)

Wilcox et al. (2004)

9

Complex spin/spinal cord anatomy

Complex spin/spinal cord anatomy

Complex spin/spinal cord anatomy

Complex spin/spinal cord anatomy

Greaves et al. (2008)

Khuyagbaatar et al. (2016)

Russell et al. (2012)

Kimpara et al. (2006)

Scifert et al. (2002)

Category

Author/year

Investigate injuries from simulated severe frontal impacts

Determine whether model-predicted strains correlate with regions of tissue damage in experimental studies

Compare regions and magnitudes of peak spinal cord stresses and strains for different spinal injury mechanisms

Compare regions and magnitudes of peak spinal cord strains for different spinal injury mechanisms

Determine the mechanism of the fracture event in spinal burst fracture

Human cervical spine and spinal cord including vertebrae, intervertebral discs (modelled as rigid bodies), white and grey matter, dura with nerve roots and CSF (including fluid-structure interaction). Rodent cervical spine (C3-C6) and spinal cord including grey and white matter, dura, CSF, intervertebral discs, spinal ligaments and dural attachments. High-speed (~1 m/s) contusion and dislocation simulations. Exquisitely detailed human head-neck complex including scalp, skull, brain, CSF, dura, vertebral bodies, intervertebral discs, muscles, ligaments, spinal cord with white and grey matter and pia,

Human cervical spine (C4–6) and spinal cord including spinal ligaments, dura and dural attachments (all modelled as linear elastic materials)

Human cervical spine (C3-C7) and spinal cord including vertebrae, intervertebral discs, spinal ligaments, spinal cord grey and white matter, dura, nerve roots and sheaths, and CSF Bovine spine and spinal cord including intervertebral disc, posterior longitudinal ligament (PLL) and dura. Detailed material properties for bone and disc

Predicted principal strains in the white matter did not exceed spinal cord yield strain measured experimentally in animal tissue, but did for grey matter. Ligament rupture predicted in some simulations because of excessive motion between the head and torso

Mechanism of injury proposed as: the PLL becomes stretched and the cord and dura become compressed by the bone fragment. These deformed before forcing the fragment back towards the vertebral body. Final position of fragment did equal maximum dynamic canal occlusion. Focal regions of peak strain were predicted for contusion and dislocation, and more diffuse strain distribution along the length of the cord for distraction, with the highest peak strains in dislocation, and the lowest for distraction Region of high stress and strain was localised anterolaterally in dislocation but posterolaterally in contusion. For distraction, the peak stress was localised to the grey matter at the lesion site, but peak strains were diffuse along and across the cord. Tissue damage was correlated with peak maximum principal strain in the white matter for both contusion and dislocation.

Significant differences in cord stresses and strains for patients with versus without neurological deficit; spinal cord stresses and strains were correlated with neurological deficit in the matched clinical cases of SCI. Impactor alignment and pre-load in experimental contusion models are strong contributors to mechanical and functional outcomes observed in in vivo experiments Higher stress in the anterior spinal cord than in the posterior spinal cord in neck flexion

Human cervical cord; subject-specific models of clinical SCI cases; cord sizing and compression magnitude/location based post-injury MRI scans; models included grey and white matter, pia, dura and denticulate ligaments, but not spinal geometry nor CSF Primate unilateral cord contusion including spinal cord, dura, CSF and a shell boundary to represent the spinal canal

Compare simulated cord stresses and strains with neurological deficits from clinical SCI cases

To assist development and refinement of an in vivo primate model while minimising animal use Determine spinal cord stresses and strains in neck flexion and extension injuries

Major findings

Methodology

Study intent

Table 3 Computational models of spinal column injury leading to SCI.

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calibrated linear sliders. CSF reservoirs were at both ends, and a water column was used to maintain physiologic hydrostatic pressure. The posterior canal was simulated with a flat plate, mounted on a load cell, and the impact was delivered via weight drop apparatus with instrumented tip. Fluid pressure was measured with miniature fibre-optic transducers. Peak CSF pressure at each location was proportional to impact velocity, and inversely proportional to the thickness of the CSF layer, while cord compression and impact load were reduced with increased CSF thickness at each impact velocity. Bilston and Thibault (1993, 1997) moulded a surrogate cervical spinal cord from silicone gel, and painted its posterior surface with a grid of enamel dots to measure cord strain via a window created by removing the laminae of a plastic reproduction of a spine (and skull). The model was tested quasi-statically to 65° flexion; the upper cervical cord moved 2.3 mm caudally during flexion, which was similar to motion observed in an in vivo MRI study (Bilston and Thibault, 1993). In a later study, the surrogate construction was modified to incorporate an outer shell of stiffer silicone gel and polyester fibres. The enamel dot grid was painted on the mid-sagittal plane of the cord and viewed via slots in the left laminae, during dynamic tests. During flexion, tensile strains were lowest in the caudal cord and highest in the middle to rostral cord, with peak strains ranging from 13% at C6-C7, to 34% at C3–4, at strain rates of up to 15 s−1. In extension tests, compressive strains of between 0 and 15% were observed, along with some local tensile strains of 10–12% (Bilston and Thibault, 1997). The results of the first four studies outlined above provide insight into the contributions of the spinal dura, ligaments and CSF to spinal cord tissue deformation. The findings help to define the anatomical features that may be important to include in cadaveric and computational models of SCI, and also the limitations of animal models (which, for example, can have substantially different CSF and dura characteristics, and lack contribution of spinal ligaments due to the injury mechanism employed). Considered together, these works illustrate the concept that CSF has a mechanically protective effect and so persons with less CSF surrounding their spinal cord would have greater risk of injury, all other things being equal; and, that the velocity and surface area of a bony fragment impinging on the cord will (in part) determine the severity of the cord deformation during the event. The unique experiments of Bilston and Thibault (1993, 1997) provide interesting data that may be of use to validate computational models of inertial loading of the neck (by the head) that seek to include measures of spinal cord stress and strain as outcome measures. The knowledge that tissue strains can be variable along the length of the cord (in an injury mechanism that does not include bony fracture or dislocation) may be interesting from a clinical perspective. However, it is important to note that this model did not include segmental nerve roots (which are expected to provide tethering of the cord), nor were the kinematics of the surrogate spinal column validated (which may influence the transmission of strain from the column to the cord).

their effects on outcomes can be determined (such as that performed by Sparrey et al. (2009) for spinal cord). Another great challenge lies in experimentally validating models of SCI; i.e. in achieving high speed loading conditions and in sourcing an appropriate experimental model (e.g. cadaveric or in vivo animal model). To this end, it would be advantageous for research groups with different resources and modelling/ experimental to collaborate and assist in cross-validation of models. Furthermore, deciding which tissues or structures to include and the extent to which geometric accuracy is important is also difficult. Two studies by Persson et al. (2011a, 2011b) have demonstrated the importance of including the fluid-structure interactions due to the presence of CSF, and the significant effects of altering the CSF thickness on predicted spinal cord deformation. To this end, it may be useful to establish a set of minimum inclusions in computational models simulating SCI. 5. Inanimate physical models used to understand mechanics of injury A number of physical models have been created, using a combination of synthetic and ex vivo materials, to better understand the mechanics of SCI (Table 4). In contrast to the cadaveric models discussed above, the majority of these models have simulated dura and cerebrospinal fluid, and have focussed on understanding the importance of these components in the mechanics of the injury event. Hall et al. (2006) created a model of cord insult in which bovine spinal cord (with and without dura) was fixed vertically in a materials testing machine. The posterior elements were simulated with simple (flat) or anatomically accurate plastic models (Table 4). Three pressure transducers were inserted transversely into the cord and a circular pellet was projected at the cord, using a calibrated pneumatic apparatus. Increased pellet velocities led to increased spinal cord occlusion and peak pressure at the impingement site. Presence of the Dacron posterior longitudinal ligament (PLL) surrogate reduced the peak pressure and cord compression, and this effect was greater at higher PLL pre-strain (8 vs. 0%). Posterior element geometry and presence of the dura had little effect on peak pressure and cord compression in this model. Jones et al. (2008) used a similar technique to establish the effect of CSF on cord compression, and to validate the dynamic behaviour of a surrogate spinal cord. The bovine cord and synthetic cord (Kroeker et al., 2009) tests, were carried out with bovine dura and CSF, with bovine dura only, and without dura. In the absence of CSF, the dura had no significant effect on cord deformation. However, spinal cord deformation was significantly reduced, although not eliminated, in the presence of CSF when compared to the bare state. The surrogate cord exhibited similar maximum deformation to the bovine cord when CSF was present, but lower deformation magnitude and duration, and lower restitution ratio (a measure of energy absorption) in the absence of CSF or dura. Persson et al. (2009) used a similar model to evaluate the effect of bone fragment size (independent of mass, and therefore impact energy) and CSF layer on cord deformation. Circular pellets were projected at bovine spinal cord–dura constructs, or a synthetic cord (based on the QMskin model, (Kroeker et al., 2009) and dura (vinyl polymer) model (Table 4). CSF reduced cord deformation regardless of pellet size, and reduced pellet size led to increased maximum cord deformation. Duration of deformation was increased in the presence of CSF, and with larger diameter pellets. The presence of dura, without CSF, did not alter the results from cord only. Jones et al. (2012a) used some similar components to the previous models (Jones et al., 2008; Persson et al., 2009), but utilised high speed x-ray to visualise deformation of the radiopaque surrogate cord (Jones et al., 2008; Kroeker et al., 2009) during the impact event. The surrogate cord and dura (polyethylene) were suspended horizontally, and secured either end such that they could be tensioned independently on

5.1. Challenges The distinct advantage of the models described above is the ability to vary physical parameters in a controlled manner, to limit or eliminate specimen variability, and potentially to visualise the impact event without a bony canal. There are a number of challenges in ensuring that these models, while simplified, still retain relevance to the in vivo situation, and while these challenges are similar to those facing computational models, they must be addressed in the physical world. Firstly, the ex vivo or synthetic materials used to model the in vivo tissues must have similar mechanical properties, which is challenging given the viscoelastic and heterogeneous nature of most biological tissues. The process of developing these physical material models, can involve a large amount of trial and error, and the rigorous mechanical testing required can be time consuming. Forming the materials into an appropriate geometry (usually highly simplified) is also challenging. 10

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Table 4 Inanimate physical models used to understand mechanisms of SCI. Author/year

Category

Study intent

Methodology

Major findings

Hall et al. (2006)

Bovine cord

Investigate effect of pellet (7.2 g, 20 mm diameter) impact speed 2.5, 5, 7.5 m/s, posterior element configuration, posterior longitudinal ligament (PLL) strain (0, 14%), on cord compression measured with high speed video.

Increased velocities: increased cord compression, peak pressure. PLL: reduced peak pressure, cord compression; greater effect with strained PLL. Posterior element geometry and dura: little effect on peak pressure and cord compression.

Jones et al. (2008)

Bovine cord/surrogate spinal cord, with bovine dura.

Investigate effect of CSF on cord deformation and restitution ratio.

Persson et al. (2009)

Bovine cord with bovine dura; surrogate cord, with synthetic dura.

Investigate effect of bone fragment size (pellet 7 g, diameter 77, 157, 314 mm2), and CSF, on cord compression.

Jones et al. (2012a)

Radiopaque surrogate cord, synthetic dura

Investigate effect of impact velocity (1.6, 2.5, 3.8 and 4.9 m/s) and CSF layer on cord compression and CSF pressure.

Bovine cord (130–180 mm length) with and without bovine dura (no CSF) fixed vertically in materials testing machine, 8% pre-strain (equivalent to in vivo neutral posture). Pellet (7.2 g, 20 mm diameter). Pressure transducers at, and ± 20 mm from, impact site. Cord compression measured with high-speed camera. Bovine/surrogate cord with and without bovine dura, with and without CSF simulant (saline), fixed vertically in materials testing machine. Cord deformation measured with high speed camera. Bovine or surrogate cord with bovine or synthetic dura, respectively, with and without CSF simulant (saline), fixed vertically in materials testing machine. Saline (0.9%) was used to simulate CSF and a flat plate formed the posterior canal. Cord deformation measured with high speed camera. Radiopaque surrogate cord, synthetic dura with CSF simulant (water) mounted horizontally. High-speed x-ray measurement of cord compression, pressure transducers at six locations.

Bilston and Thibault (1993)

Surrogate cord tethered in plastic cervical spine/ skull

Investigate cord motion within model at 65° flexion (quasistatic)

Bilston and Thibault (1997)

Surrogate cord tethered in plastic cervical spine/ skull

Investigate cord strains during dynamic tests ( ± 65° from neutral, 28–40 g peak acceleration, 50–65 ms pulse duration).

Surrogate cord of silicone gel, with enamel dot grid painted exteriorly. The gel material had similar tensile properties (at loading rates 0.02–0.3/s) to human spinal cord. Surrogate cord of silicone gel and 40 axial polyester fibres, with enamel grid in mid-sagittal plane. Plastic spine/skull with visualisation windows through laminae.

CSF: reduced cord compression; dura without CSF: no effect on cord compression. Surrogate cord: similar cord compression with CSF/dura, but lower deformation without CSF or dura. CSF: reduced cord deformation regardless of pellet size, increased duration; Reduced pellet size: increased cord deformation and decreased duration.

Peak CSF pressure: proportional to impact velocity, inversely proportional to CSF layer thickness. Cord compression and impact load reduced with increased CSF thickness at each impact velocity. Upper cord moved 2.3 mm caudally, similar to motion observed in vivo with magnetic resonance imaging.

Flexion: tensile strains lowest in caudal region, highest in middle to rostral; peak strains 13% (C6C7) to 34% (C3–4); strain rates up to 15 s−1. Extension: compressive strains 0–15%; tensile strains 10–12%.

combinations thereof. Each of these approaches has distinct advantages and disadvantages, and each will benefit from incremental improvements in our understanding of the biomechanical system and advances in technology. The high-speed nature of these injuries leads to challenges with the sensors, actuators and instrumentation required to produce and interrogate these models. Advances in high speed videography, high speed simultaneous data acquisition, and actuators and their control systems in recent decades, have improved our ability to physically model and measure simulated injuries occurring at nearrealistic speeds. Similarly, advances in computational power and modelling software have enabled more complex modelling of nonlinear viscoelastic tissues, fluid-solid interactions, and higher speed interactions in FE simulations. Important lessons that have been learned from the models thus far include: tissue injury and functional deficit are dependent on injury mode or mechanism and magnitude, post-injury imaging may not represent the “worst case” impingement of the spinal cord, the CSF reduces deformation of the cord during a mechanical insult, various tissues of the spine and spinal cord have different injury thresholds so injury criteria need to consider this accordingly, and evoked potentials and imaging show potential in being able to monitor for and predict neurological deficit in in vivo animal models.

Secondly, realistic boundary conditions must be provided for both the solid and fluid materials; for example, the spinal cord and dura exist in a state of tension or pre-strain in vivo, the CSF has a baseline and pulsatile pressure, and the spinal canal is semi-rigid and provides the surfaces which contact the cord-dura construct during the injury event. Each of the models created has selected particular boundary conditions to simulate, but none provide all, and this limits their utility to gaining a broad understanding of the mechanical effect of the parameter being studied. 6. Discussion and conclusion Devising methods and techniques to study the complex interaction between the spinal canal and its contents during injury events, and the implications of that interaction on risk and characteristics of neurological damage, is extremely challenging. The greatest challenge lies in studying an injury as it occurs - it is impossible to produce simulated injuries in living humans (an issue that is not common to the study of other disease states), hence the need for modelling approaches. This review has described a wide variety of engineering approaches that have been taken to improve understanding of the biomechanics of SCI, spanning cadaveric, computational, animal and physical modelling, and 11

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The most important issues in the field of SCI that can be advanced through these engineering approaches are; preventing spinal injuries, and advancing and refining the models that are used to discover and test therapies to treat SCI. Refining each of the modelling approaches described in this review will improve the injury criteria that are used to assess injury prevention strategies, and improve the management of spinal trauma which may, in turn, reduce or prevent additional SCIs. Cadaveric models of traumatic spinal column injury (with measures of cord compression) are the most likely modelling modality to lead to improved knowledge of injury mechanics, and therefore improved injury criteria and anthropometric test devices, both of which are essential for assessment of prevention devices and strategies. An effort should be made to include dynamic measurement of spinal cord compression or canal occlusion in all cadaveric studies of spinal column trauma. If this data is shared amongst groups, this will build a dataset suitable for creating new injury risk criteria that consider the spinal cord across all column injury mechanisms. Further, to improve FE models of spinal trauma, cadaveric studies should include outcome measures useful for validation of said models, such as bony and ligamentous surface strains, intervertebral kinematics, and bony deformations. These data should be made available by experimentalists to FE modellers. The most important advancement for in vivo animal modelling of SCI will be increasing the level of collaboration between engineers, clinicians and scientists. Increasing the use of models with different injury mechanisms (which lead to different functional deficits and pathologies), including closed-column models such as dislocation and distraction models, should improve clinical translation of SCI therapies. The obvious challenge with this vision is that these models are more technically difficult to use than the standard contusion models, and can require post-injury spinal stabilisation devices. Increasing interdisciplinary collaboration would improve the uptake and utilisation of technically challenging (but more clinically relevant) SCI models. The most important outcomes from computational modelling will come from open sharing of resources and increased global cross-disciplinary collaboration. It is clear that in several research teams globally, the resources exist to develop detailed and realistic models to simulate spine/spinal cord injuries, and to use these to answer clinically and scientifically pressing questions. Presently, teams work in relative isolation using customised models to address specific questions. The computational models, with considerable development time and cost, are then often not used in subsequent studies or to address additional aims. Increased sharing of model resources and increased global, crossdisciplinary collaboration, will advance injury prevention strategies, improve injury criteria, and improve diagnostic and prognostic capabilities. The continued application of engineering fundamentals and contemporary technology in SCI biomechanics research will lead to improved understanding of the mechanical context in which the biological injury occurs, and insight into the complex interaction between the external loads applied to the body and the resultant interaction between the spinal column and cord. It is likely that this will incrementally lead to improved injury criteria for use in crashworthiness assessments with anthropometric test devices, improved multibody dynamic reconstructions and injury prediction models, and improved animal models that are able to more accurately screen surgical, medical and pharmacological treatments for efficacy and thereby increase the success of clinical translation.

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Conflicts of interest None. Acknowledgements CFJ gratefully acknowledges salary funding through an Early Career 12

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