Jump to content

Documentation:FIB book/The Relative Contributions of Rotational Loading and Accelerations in the Injury Mechanisms of Traumatic Brain Injury

From UBC Wiki

Overview

Introduction

Epidemiology

Traumatic Brain Injuries (TBI) occur when external forces, including impact or inertial loading, result in movement of the brain within the skull, often leading to structural and functional damages or impairments[1] [2]. This can lead to varying degrees of brain damage, ranging from mild concussions to severe cognitive impairment. TBI is a global health problem, with 2.8 million people sustaining a TBI in the United States alone, including an estimated 50,000 fatalities [1]. Beyond the immediate medical consequences, TBI is one of the leading causes of disability and affects an estimated 3.2 to 5.3 million individuals in the United States[1] [3]. Those surviving TBI often experience long-term cognitive, psychological, and emotional impairments, impacting their quality of life and ability to return to work. TBI also poses a massive economic and healthcare burden, incurring $60 billion USD in medical costs and contributing to 30.5% of injury-related fatalities in the United States [1] [3].

However, despite their high incidence, the underlying mechanisms of TBI remain an understudied phenomenon. Two key contributors to TBI damages are translational and rotational acceleration, but the relative contributions of these motions are largely unknown. Translational acceleration refers to linear motion and has typically been shown to be implicated in skull fractures and direct impacts or contusions. Conversely, rotational acceleration, which contributes to angular or rotational motion, is typically associated with diffuse axonal injuries and subdural hematomas. However, while modern advancements have cited rotational forces in addition to linear motions as a major contributor to TBI, advances in rotational testing methods have been developed, and many testing standards have not kept pace, revealing discrepancies in understanding TBI mechanisms. Proper comprehension of these mechanisms is critical for developing robust safety equipment, maintaining adequate regulations, and advancing modern clinical understanding.

This literature review provides insights into the mechanisms of traumatic head injury and the respective contributions of rotational loading and accelerations.

Long-Term Effects

The long-term effects of TBI can range in severity and duration, depending on both the mechanism of injury and injurious outcomes. Common impairments due to TBI include cognitive, neurological, psychological, and physical effects [4][5].

  • Cognitive effects of TBI may include memory loss, cognitive decline, and learning difficulties. These can affect a person's ability to concentrate, communicate, problem solve, retain information, and interact with others.
  • Physical effects include chronic pain, persistent headaches, sensory issues, motor problems, paralysis, and fatigue.
  • Neurological effects include seizures and brain atrophy, which could lead to neurodegenerative diseases such as Alzheimer’s disease or Parkinson’s disease.
  • Psychological effects include mood changes, anxiety, depression, and behavioral issues.

These consequential impacts of TBIs often significantly impact the individuals' quality of life, decreasing their ability to maintain social interactions and communication. This may lead to adverse challenges such as unemployment, illicit drug use, and ongoing medical support [4]. It is important to note that these effects' occurrence, severity, and duration depend on factors such as age, mechanism of injury, and treatment.

Injury Mechanisms

Rotational forces or combinations of rotational and linear forces are more common and are understood to cause more TBIs. Models show that 90% of the strain in brain tissue during head acceleration is due to rotational kinematics [4]. These include oblique impacts, where the angle of impact is most commonly in the range of 30-40 degrees from the ground, and tangential impacts, which occur around the head’s center of gravity [6].

Rotational forces on the brain lead to shear stress and strain within the brain tissue, which causes diffuse injuries throughout larger brain areas than the focal injuries caused by linear impacts.

These rotational brain injuries include:

  • Diffuse Axonal Injury (DAI): Caused by large amounts of shear and strain from rapid rotational acceleration in the brain, leading to axon damage. This can lead to loss of consciousness, coma, memory loss, and motor deficits [6].
  • Subdural Hematomas: Caused by rotational forces rupturing veins between the brain and the dura, forming blood clots [6].
  • Concussions: Being the most common head injury, brain lesions are caused by both linear and rotational accelerations but are significantly more likely to occur during rotation due to strain within the brain [6].
  • Contusions: Can be caused by skull fracture or by shearing of the brain tissue against the inner surface of the skull and dura, which can cause hemorrhage, edemas, and more severe brain injuries [6].
  • Neuroinflammation: The rotational forces of the brain within the skull can cause tiny blood vessels to tear and lead to brain inflammation, which can disrupt nerve connectivity [7].

The image in Figure 1 demonstrates the difference between a purely linear acceleration impact and an oblique angle impact and the types of injuries that these injury mechanisms can cause.

Figure 1. Comparison of the biomechanics of a linear acceleration impact (upper) and an oblique angle impact (lower)

Common Injury Modes

Understanding the common causes of TBIs is essential for understanding the injury biomechanics, prevention mechanisms, and previous studies performed on brain injuries. Many scenarios lead to TBI, with the most frequent listed below [2]:

  • Falls are the leading cause of TBI, primarily among seniors and young children.
  • Sports Injuries are very common among adolescents and occur often in high-contact sports. Football tackles, hockey body checks, boxing matches, and soccer headers are some of the most common causes of sports-related injuries.
  • Vehicle collisions have a high incidence of TBI, and the type of TBI varies based on the collision type. Many collisions cause whiplash, which may lead to TBI; collisions involving motorcycles, bicycles, and pedestrians are also common causes of trauma to the brain.
  • Explosive blasts put military personnel at an elevated risk of TBI due to the pressure waves that can impact the brain during explosions, as well as other combat-related injuries.

These causes of TBI involve complex biomechanics and typically involve a combination of linear and rotational forces. In real-world injuries, head impacts are very often off-center from the brain's center of gravity (tangential/oblique impacts) or are due to the brain shearing within the skull during acceleration forces (whiplash). Because of this, it is essential to have a strong understanding of the relationship between rotational acceleration and TBI.

Existing Literature

Foundational Studies on Rotational Acceleration and TBI

Modern understandings of rotational acceleration as a contributor to TBI are based upon foundational proposals of kinematic mechanisms of brain injury. Holbourn (1943) was one of the foundational researchers who proposed that rotational acceleration could contribute to TBI, hypothesizing that rotation from blunt TBI caused injury, while translation played an insignificant role [8]. This was a substantial advancement in the field, as historical understandings of TBI cited linear forces as the major contributor and only factor to injury.

Holborn used theoretical models and mathematical derivations to argue the behavior of the skull and brain during impacts, deriving that brain injury concludes when its constituent particles are displaced to the extent where realignment does not occur [8]. This is characterized by shear strains resulting from rotational accelerations, rather than compressive or rarefaction strains of linear forces, thus leading to Holborn's conclusion that shear strain is the rate of measure for probability of brain injury [8]). Moreover, the brain's low shear modulus and relatively high bulk modulus indicate greater deformation when subjected to rotational accelerations. Holborn additionally cited Grundfests’ (1936) work with peripheral nerves, which found that nerves would continue to conduct with hydrostatic pressures of 10,000 lb/sq in applied, though minor applications of pressure in various directions (thus replicating shear strains) was sufficient for nerve dysfunction [8]. Holborn extended Newton’s laws of motion to TBI, concluding that its severity in impact is proportional to the total change of velocity rather than the accelerative rates present [8]. Holbourn’s work provided a theoretical foundation for later experimental studies, which validated his hypothesis.

Building off of Holbourns theorems, Gennarelli et al. (1990) demonstrated that rotational acceleration is a primary factor in producing diffuse axonal injury (DAI) in work with primates. DAI is a diffuse type of TBI and is believed to be a primary injury type, occurring at the time of the accident rather than after the initial trauma [9]. As such, the study of DAI is highly relevant to deriving understandings of TBI during impact. In this work, primates were subjected to controlled angular accelerations in sagittal, coronal, and oblique planes.

One of the most significant findings from Gennarelli’s work at the time was the identification of a direct relationship between rotational acceleration and unconsciousness. It was found that head movements with purely linear acceleration did not result in traumatic unconsciousness, yet the introduction of rotational components drastically increased the likelihood of unconsciousness as an outcome [9]. This study validated the hypothesis that the shear strain induced by rotational forces is a major contributor to neurological impairment, particularly traumatic unconsciousness [9].

Finding the incidence of traumatic unconsciousness greatest in purely coronal accelerations, Gennarelli expanded his work to construct physical models of the baboon skull and brain structure, which permitted measurement of distortions within the surrogate brain and facilitated prediction of human injury thresholds through empirical scaling techniques [9]. Applying only coronal accelerations, and using the same load levels and acceleratory ranges, Gennarelli was able to use the physical model to reduce the number of degrees of freedom in load configuration, permitting the use of peak angular deceleration magnitude as a single independent loading variable, thus providing differentiations between mild and severe TBI outcomes, finding that mild DAI, rather than concussions, occurred with higher levels of angular accelerations [9]. Gennarelli was able to conclude that shear strains and brain tissue pathology spatial distributions mirrored one another, with brain regions that are most often associated with DAI pathology associated with the highest strain levels in the physical models [9]. This finding cemented shear strain as an index for primary tissue injury due to rotational loads, empirically deriving rotational acceleration thresholds for DAI in humans, which were later used as a reference in future injury threshold evaluations.

These studies established that rotational motion alone, without direct head impact, could induce severe brain injuries, including widespread axonal damage, hemorrhages, and coma-like states, with shear strain emerging as a key determinant of injury severity and neurological impairment in TBI.

Recent Testing Methods and Tool

In Vivo Studies

In vivo models are extremely important in the research of rotational TBI as they allow for the evaluation of cognitive and behavioural effects as well as secondary injuries such as inflammation. In vivo TBI models can be broadly classified into open-head and closed-head injury (CHI) models.

Open-Head Models

Open-head injury models involve the direct application of mechanical force to the exposed dura mater via craniotomy. As the mechanical force is applied directly to the brain, and there is no head movement, these models are predominantly used to replicate injuries caused by linear acceleration and focal injury [10].  Two widely used open-head models are the fluid percussion (FP) and controlled cortical impact (CCI) models [11][12].

Closed-Head Models

Closed-head injury (CHI) models are important tools for studying rotational acceleration and its effect on TBI. These models involve an indirect transmission of force to the brain, which can be either induced by a direct impact or an indirect form of inertial loading [10]. As the vast majority of human TBI occurs without skull fracture, CHI provides a more relevant mechanism than its open-head counterpart [13]. In addition to this, the non-surgical nature of the models allows for multiple impacts to be studied, as well as the long-term consequences to be observed with a simplified clinical setup.

A widely recognized CHI model is the Closed-Head Impact Model of Engineered Rotational Acceleration (CHIMERA), developed by Namjoshi et al. This is a non-surgical model that provides a controlled impact to rodents' heads to induce rotational TBI. CHIMERA allows for precise control and semi-automation of impact parameters, including velocity, angle, and location, which ensures repeatable methods across studies. It was proven to replicate fundamental mechanisms and responses of impact TBI in humans, including axonal damage, tau phosphorylation, and inflammatory reactions [14]. 

  • Figure 2. Schematics of the CHIMERA device (A) Provides an overview of the parts of the device
    Figure 2. Schematics of the CHIMERA device (A) Provides an overview of the parts of the device
  • (C) Demonstrates the location of the impactor on the mouse head, including the location of the brain.
    (C) Demonstrates the location of the impactor on the mouse head, including the location of the brain.

Another category of CHI models focuses on simulating blast-induced TBI that occur from explosions and are commonly seen in military personnel. A study by Bugay et al. used air-driven pressure pulses to subject mice to repeated blast scenarios; further, work by Rubovitch et. al used real explosives to induce blast-related TBI. Both studies showed effects consistent with those seen in humans, including decreases in cognitive and behavioural test scores and seizures [15][16].

Finally, to investigate pure inertial loading, several studies have focused on reproducing rapid head motion without an impact, simulating conditions including whiplash-induced TBI. In these studies, large animal models, primarily pigs, are frequently used due to their anatomical and biomechanical similarities to the human brain. Notably, the porcine-HYGE model by Cullen et al.. is a widely recognized model that is defined to be biomechanically representative of human rotational TBI. This model uses a HYGE pneumatic device and a custom linkage system to deliver non-impact rotational acceleration to pigs’ heads. It is shown to reproduce the diffuse axonal injury, neuroinflammation, and prolonged unconsciousness observed in severe human TBI [17].  

Animal Model Biofidelity

In terms of biofidelity, large animal models are preferable to rodent models due to various parameters, including their mass. Further, non-human primates, pigs, and dogs have brain geometry with folds and grooves (gyrencephalic), similar to a human brain, whereas rodents have a smooth (lissencephalic) brain geometry. The presence of these folds and grooves, or sulci and gyri, results in significantly more brain deformation during impact and angular acceleration compared to smooth brains [18]. In addition to this, these large animals have a white matter to gray matter ratio that is more representative of the human brain [15][16][19].

Computational Simulations and Finite Element Analysis (FEA)

In addition to the in-vivo testing methods discussed above, computational simulation tools, specifically finite element head models (FEHM), are an increasingly relevant tool for studying TBI. SIMon, a simulated injury monitor system FEHM is a software package which has recently been developed to advance the interpretation of TBI mechanisms. Using detailed geometric specifications and anatomical characteristics, the SIMon FEHM can been used to calculate three primary injury metrics: Cumulative strain Damage Measure (CSDM), Dilational Damage Measure (DMM), and Relative Motion Damage Measure (RMDM) from input kinematic data of crashes tests, sports impacts, or other user-specified scenarios [20]. CSDM correlates for DAI, DMM estimating contusion probability, and RMDM correlates for acute subdural hematomas. The extensive coverage and variation in TBI severity SIMon provide a dependable and highly applicable model for TBI study. SIMon FEHM has been further validated against cadaveric data and animal brain injury experiments, thus recognized as a robust and well-defined model.

Injury Thresholds and Criteria

Historically, head injury assessment has relied on criteria such as the Gadd Severity Index (GSI) and the Head Injury Criterion (HIC) [21].  These injury criteria are discussed in more depth in the Injury Criterion sections of this Wiki. These metrics are based primarily on linear acceleration parameters and do not capture the effects of rotational acceleration. This gap in injury criteria metrics has led to the development of the Brain Injury Criterion (BrIC), a rotational-specific injury criteria.

Brain Injury Criterion

The BrIC was developed by researchers at the National Highway Traffic Safety Administration (NHTSA) in collaboration with Bowhead Systems Management, Inc. and Virginia Tech, to address rotational acceleration in brain injury criteria. The BrIC model was derived from the SIMon model, as previously discussed. Further, it used scaled animal model data to derive strain and stress-based injury thresholds. Finally, 50th percentile HIII, ES-2re, and WorldSID ATDs were used to establish the kinematic parameters of rotational acceleration and velocities from test data. The authors suggest that BrIC be used in conjunction with HIC to provide a more comprehensive assessment of brain injury risk [22].

Discussion

Controversies

Animal models have been widely used in TBI research to study the effects of rotational acceleration. One of the most well-known and earliest studies was conducted by Gennarelli et al., who used primates to demonstrate that rotational forces could cause diffuse axonal injury, a type of brain injury associated with TBI [9]. This model was designed to produce coronal rotational acceleration, where researchers found that this motion resulted in the most widespread axonal damage in primates [23]. While these studies provided valuable insights in understanding rotational forces in TBI, they also highlighted the challenges of translating findings from animal subjects to humans [9]. Anatomical differences between animal and human brains, such as skull shape, brain size, and white-to-gray matter ratios, affect how injuries develop [19]. Similarly, rodent models are also commonly used in TBI research, but one of the major challenges associated with this is the structural differences between rodent and human brains, particularly skull shape and impact location [24]. Human TBIs often result from forces striking the frontal, occipital, or temporal regions of the head, where the brain is damaged at both the impact site and its opposite site due to rebounding forces [24]. In contrast, rodent models are subject to force at the top of the skull, which does not represent how injuries occur in humans [19][24]. Studies have attempted to improve rodent models by introducing frontal impacts with rotational acceleration that produce more widespread brain damage; however, even with these modifications, rodents fail to capture the exact injury patterns seen in humans [23]. Furthermore, biological differences between animals and humans can limit the reliability of results, raising questions about the accuracy and necessity of using them in brain injury research. This also raises ethical concerns regarding the welfare and treatment of animals as they likely endure pain, distress, or long-term harm while testing.

Limitations

One of the most significant limitations in the majority of studies investigating TBI injury mechanisms is lack of consistency and industry standards in defining impact conditions and experimental protocols which effectively distinguish between linear and rotational acceleration [25]. A persistent challenge in experimental design has been difficulty in isolating the effects of linear accelerations during testing, largely due to the inherently complexities of controlling head movements and impact conditions [25]. As such, there have been few advancements in successfully decoupling linear and rotational components, limiting the precision and effectiveness of injury mechanism analyses. In response, researchers have proposed more comprehensive approaches in quantifying head injury risks using both linear and rotational accelerations; for example, one study discusses a combined probability model, which evaluates injury severity with both acceleration considerations, demonstrating an improved predictive capability compared to models relying on a single kinematic parameter [26]. However, this same study also reported inconsistencies in the predictive power of linear and rotational accelerations across injury scenarios, emphasizing the ongoing uncertainty regarding their respective roles in TBI [26]. Although Holbourn and Gennarelli et al. found rotational accelerations to be strongly implicated in TBI injury mechanisms, the relative contribution with linear accelerations remains adequately quantified. This highlights the need for further experimental research employing advanced impact models capable of studying linear and rotational forces in tandem, and using impact methods which can isolate and analyze these respective influences.

A method which has been traditionally used to study these relative contributions, are head impact sensors [27]. Head impact sensors are quite advantageous for the field of head biomechanics, largely due to their ability to provide real-time monitoring for immediate data collection, non-invasiveness, and role in correlating kinematic data with injury risks [27]. Within the field of TBI, their strongest feature pertains to an ability to measure both linear and rotational accelerations during impacts, directly addressing the current gap in literature [27]. However, while a valuable tool in offering real-time kinematic data, the utility of head impacts sensors are inherently limited by their ability to accurately detect and characterize a range of impact conditions [25]. As emphasized by Wu et al., many modern head impact sensor systems exhibited reduced sensitivities to low impact conditions, potentially leading to underestimations of injury risk. While missing low-magnitude impacts may be less critical in studies focused solely on severe head impacts, this limitation becomes significant in TBI- one of the most common rotational injury outcomes, concussions, have been reported at acceleration levels as low as 29.3g [25]. As a result, the inability of certain head impact sensors to reliably detect such impacts raises several concerns, especially given the majority of real-world TBI impact methods are of mild to moderate severity (ie. falls, sports injuries, as detailed in 1.4). Improving head impact sensor reliability across the full spectrum of impact magnitudes would ensure reliable data collection, strengthen injury prediction models, and ultimately support the development of accurate quantifications of the contributions of linear and rotational accelerations in TBI. Additionally, physiological tolerance patterns play a role in successfully identifying, and quantifying levels of rotational forces which have injurious implications. For example, the brain's tolerance to high rotational accelerations decreases linearly with prolonged exposure- the timing and duration of data collection thus critically influences the representativeness and accuracy of recorded kinematic data [28]. Together, these technical and physiological limitations underscore the need for the development of improved head impact sensor systems, and more comprehensive injury models capable of capturing the full complexity of TBI biomechanics.

Future Work

Despite significant advancements in understanding the role of rotational acceleration in TBI, several gaps continue to remain in both biomechanical understandings of rotational acceleration injury contributions, and appropriate experimental methods to study it. A substantial portion of research thus far has relied on animal models and computation simulations, which, while valuable, often fall short in replicating the anatomical complexities and biomechanical variability of real-world human scenarios. Diversity in patient populations - including factors such as age, sex, ethnicity and pre-existing health conditions- can further influence the accurate modeling of TBI, underscoring a need for increased representative data and methods in TBI clinical research. Moreover, inconsistencies in experimental setups, and a lack of standardized protocols for inducing and measuring axis-specific accelerations have hindered cross-study comparison validity and broader variations of findings.

A critical avenue for future work in response to TBI lies in the application of biomechanical data towards improving helmet design. Current helmet designs have been optimized to mitigate linear accelerations, but not sufficiently address the rotational forces that have been strongly implicated in TBI injury [29]. In particular, while structure and impact absorption designs have continued to progress,  most regulatory standards for helmets still rely on linear acceleration metrics, neglecting the rotational dynamics that are increasingly recognized as foundational to TBI risk [29]. The incorporation of rotational metrics and focused injury criteria, such as the brIC, alongside would provide a more holistic assessment of helmet performance and drive effective injury mitigation strategies.  

Future studies should also prioritize experimental conditions that more realistically simulate human head motion. This includes refining injury thresholds, the development of surrogate models that better represent human brain biofidelity, and improvements in the performance of head impact sensor accuracy. As discussed, head impact sensor systems often fail to detect low-magnitude impacts; future efforts should focus on increasing their sensitivity towards low-force impacts, as many head injury mechanisms, including concussions, occur at acceleration levels below the threshold of current head sensor based devices [25]. Potential methods include optimization of signal to noise ratio (SNR), and research developments into low-noise accelerometer technology would increase the reliability of in-field data collection of head impact sensors to mild impact scenarios, thereby contributing more robustly to injury prediction models.

References

  1. ↑ 1.0 1.1 1.2 1.3 Guan, Bin; Anderson, David B; Chen, Lingxiao; Feng, Shiqing; Zhou, Hengxing (October 2023). "Global, regional and national burden of traumatic brain injury and spinal cord injury, 1990–2019: a systematic analysis for the Global Burden of Disease Study 2019". BMJ Open. 13.
  2. ↑ 2.0 2.1 National Institute of Neurological Disorders and Stroke (October 2024). "Traumatic Brain Injury (TBI)". National Institute of Neurological Disorders and Stroke.
  3. ↑ 3.0 3.1 Lang, Ji; Wu, Qianhong (February 2021). "Modeling of the transient cerebrospinal fluid flow under external impacts". European Journal of Mechanics - B/Fluids. 87: 171–179.
  4. ↑ 4.0 4.1 4.2 Bramlett, Helen M; Dietrich, W Dalton (December 2015). "Long-Term Consequences of Traumatic Brain Injury: Current Status of Potential Mechanisms of Injury and Neurological Outcomes". Journal of Neurotrauma. 32: 1834–1848.
  5. ↑ Brain Injury Association of America (March 2025). "Long-Term Effects of Brain Injury". BIAUSA.
  6. ↑ 6.0 6.1 6.2 6.3 6.4 Kleiven, Svein (November 2013). "Why Most Traumatic Brain Injuries are Not Caused by Linear Acceleration but Skull Fractures are". Front. Bioeng. Biotechnol. 1.
  7. ↑ Umfress, Alan; Chakraborti, Ayanabha; Devi, Suma P S (February 2023). "Cdk5 mediates rotational force-induced brain injury". Scientific Reports. 13.
  8. ↑ 8.0 8.1 8.2 8.3 8.4 Holbourn, A H S (October 1943). "Mechanics of Head Injuries". The Lancet. 242: 438–441.
  9. ↑ 9.0 9.1 9.2 9.3 9.4 9.5 9.6 9.7 Gennarelli, Thomas A; Thibault, Lawrence E; Adams, J Hume (December 1982). "Diffuse axonal injury and traumatic coma in the primate". Annals of Neurology. 12: 564–574.
  10. ↑ 10.0 10.1 Namjoshi, Dhananjay R; Good, Craig; Cheng, Wai Heng; Panenka, William; Richards, Darrin; Cripton, Peter A (November 2013). "Towards clinical management of traumatic brain injury: a review of models and mechanisms from a biomechanical perspective". Disease Models and Mechanisms. 6: 1325–1338.
  11. ↑ Kabadi, Shruti V; Hilton, Genell D; Stoica, Bogdan A (August 2010). "Fluid-percussion–induced traumatic brain injury model in rats". Nature protocols. 5: 1552–1563.
  12. ↑ Romine, Jennifer; Gao, Xiang; Chen, Jinhui (August 2014). "Controlled cortical impact model for traumatic brain injury". J Vis Exp. 5.
  13. ↑ Bales, James W; Bonow, Robert H; Ellenbogen, Richard G (2018). "25 - Closed Head Injury". Principles of Neurological Surgery (Fourth Edition): 366–389.
  14. ↑ Namjoshi, Dhananjay R; Cheng, Wai Heng; McInnes, Kurt A (December 2014). "Merging pathology with biomechanics using CHIMERA (Closed-Head Impact Model of Engineered Rotational Acceleration): a novel, surgery-free model of traumatic brain injury". Molecular Neurodegeneration. 9.
  15. ↑ 15.0 15.1 Bugay, Vladislav; Bozdemir, Eda; Vigil, Fabio A (January 2020). "A Mouse Model of Repetitive Blast Traumatic Brain Injury Reveals Post-Trauma Seizures and Increased Neuronal Excitability". Journal of Neurotrauma. 37: 248–261.
  16. ↑ 16.0 16.1 Rubovitch, Vardit; Ten-Bosch, Meital; Zohar, Ofer (December 2011). "A mouse model of blast-induced mild traumatic brain injury". Experimental Neurology. 232: 280–289.
  17. ↑ Cullen, D Kacy; Harris, James P; Browne, Kevin D (2016). "A Porcine Model of Traumatic Brain Injury via Head Rotational Acceleration". Injury Models of the Central Nervous System. 1462: 289–324.
  18. ↑ Finnie, J W (July 2012). "Comparative approach to understanding traumatic injury in the immature, postnatal brain of domestic animals". Australian Veterinary Journal. 90: 301–307.
  19. ↑ 19.0 19.1 19.2 Xiong, Ye; Mahmood, Asim; Chopp, Michael (January 2013). "Animal models of traumatic brain injury". Nature Reviews Neuroscience. 14: 128–142.
  20. ↑ Takhounts, Erik G; Eppinger, Rolf H; Campbell, J Quinn (October 2003). "On the Development of the SIMon Finite Element Head Model". 47th Stapp Car Crash Conference.
  21. ↑ Gao, Dalong; Wampler, Charles (October 2010). "Head Injury Criterion". IEEE Robotics & Automation Magazine. 16: 71–74.
  22. ↑ Takhounts, Erik; Ridella, Stephen; Rowson, Steve (January 2011). "Kinematic rotational brain injury criterion (BRIC)" (PDF). NHTSA.
  23. ↑ 23.0 23.1 Kilbourne, Michael; Kuehn, Reed; Tosun, Cigdem (December 2009). "Novel model of frontal impact closed head injury in the rat". Journal of Neurotrauma. 26: 2233–2243.
  24. ↑ 24.0 24.1 24.2 Zhang, Yi Ping; Cai, Jun; Shields, Lisa B E (February 2014). "Traumatic brain injury using mouse models". Translational Stroke Research. 5: 454–471.
  25. ↑ 25.0 25.1 25.2 25.3 25.4 Wang, Timothy; Kenny, Rebecca; Wu, Lyndia C (October 2021). "Head Impact Sensor Triggering Bias Introduced by Linear Acceleration Thresholding". Annals of Biomedical Engineering. 49: 3189–3199.
  26. ↑ 26.0 26.1 Rowson, Steven; Duma, Stefan M (January 2013). "Brain Injury Prediction: Assessing the Combined Probability of Concussion Using Linear and Rotational Head Acceleration". Annals of Biomedical Engineering. 41: 873–882.
  27. ↑ 27.0 27.1 27.2 O'Connor, Kathryn L; Rowson, Steven; Duma, Stefan M; Broglio, Steven P (March 2017). "Head-Impact–Measurement Devices: A Systematic Review". Journal of Athletic Training. 52: 206–227.
  28. ↑ Rowson, Steven; Duma, Stefan M; Beckwith, Jonathan G (October 2011). "Rotational Head Kinematics in Football Impacts: An Injury Risk Function for Concussion". Annals of Biomedical Engineering. 40: 1–13.
  29. ↑ 29.0 29.1 Kis, M; Saunders, F; ten Hove, M W; Leslie, J R (2004). "Rotational Acceleration Measurements – Evaluating Helmet Protection" (PDF). Can. J. Neurol. Sci. 31: 499–503. line feed character in |title= at position 39 (help)

External Links

Retrieved from "https://wiki.ubc.ca/index.php?title=Documentation:FIB_book/The_Relative_Contributions_of_Rotational_Loading_and_Accelerations_in_the_Injury_Mechanisms_of_Traumatic_Brain_Injury&oldid=860680"