Brain injury expert Dr. John Lloyd has served attorneys nationwide for 25+ years in biomechanics, human factors, helmet testing and motorcycle accident expert
Dr. John Lloyd, Research Director of Brains, Inc. announced today that football head injuries and concussions can be reduced up to 50 percent with their new helmet technology.
football helmet prototype
Tampa, FL – John Lloyd PhD, Research Director of Brains, Inc. announced their latest breakthrough in football helmet safety today. The unique new helmet technology promises to provide up to 50 percent more protection against football head injuries and concussions. The helmet technology has wide application and can be used in every kind of helmet from baby helmets to military helmets, and for all athletes at risk of concussion and head injuries such as football players, cyclists, skiers, snowboarders, skateboarders, hockey players, baseball players, lacrosse players, boxers, soccer players, equestrian / horse-riding sports, such as polo and horse racing, as well as motorcycle and race car drivers.
Recent medical research documents found that concussions and cumulative head impacts can lead to lifelong neurological consequences such as chronic traumatic encephalopathy, a degenerative brain disease known as CTE and early Alzheimer’s.
The U.S. Centers for Disease Control and Prevention, estimates 1.6 – 3.8 million sport-related brain injuries annually in the United States. Of these 300,000 are attributed to youth football players, some of whom die from their injuries every year – a tragedy difficult for their mothers and families to recover from. The severity of the issue touching both the nation’s youth and professional athletes has led to thousands of lawsuits and Congressional Hearings. Growing concern has spread to the White House where President Obama recently spoke at the Healthy Kids and Safe Sports Concussion Summit.
The BRAINS research team, led by renowned brain injury expert, Dr. John Lloyd, has worked for years on their project to help make sports safer. A controversial subject, some opponents have stated that concussion prevention is impossible. Dedicated to saving lives and preserving brain health, Dr. Lloyd and team persevered with their work leading to this new innovation. “Our results show that forces associated with concussion and brain injury are reduced more than 50% compared to similar testing with a leading football helmet,” said Dr. John Lloyd, Research Director. Results of our prototype helmet technology compared to the Riddell Revolution Speed varsity helmet are presented below: “The patent-pending matrix of non-Newtonian materials will not only benefit football, but can be utilized in all sports helmets as well as military, motorcycle and even baby helmets to improve protection and dramatically reduce the risk of brain injuries,” reported Dr. Lloyd. The materials are inexpensive, and produce a helmet that is considerably lighter and more comfortable than a traditional helmet. Two additional applications of this new safety technology include medical flooring especially in hospitals and nursing homes or child play areas , as well as vehicle interiors.
Testing Methods: A modification to the NOCSAE standard test apparatus has been developed and validated for impact testing of protective headwear to include measurement of both linear and angular kinematics . This apparatus consists of a twin wire fall test system equipped with a drop arm that incorporates a 50th percentile Hybrid III head and neck assembly from HumaneticsATD. The aluminum flyarm runs on Teflon sleeves through parallel braided stainless steel wires, which are attached to mounting points in the building structure and anchored into the concrete foundation. The anvil onto which the head drop systems impacts consists of a 350mm x 350mm steel based plate. Both the Riddell Revolution Speed varsity football helmet and prototype helmet were dropped from a height 2.0 meters onto a flat steel anvil, in accordance with ASTM standards, generating an impact velocity of 6.2 m/s (13.9 mph). The following slow motion videos show testing on an unhelmeted head and prototype using this apparatus
Instrumentation: A triaxial accelerometer from PCB Piezotronics (Depew, NY) and three DTS-ARS Pro 18k angular rate sensors (Diversified Technical Systems, Seal Beach, CA) affixed to a triaxial block were installed at the center of mass of the Hybrid III head form (HumaneticsATD, Plymouth, MI). Data from the accelerometer and angular rate sensors were acquired using National Instruments (Austin, TX) compact DAQ hardware.
Analysis: In accordance with SAE J211, data from the analog sensors were acquired at 10,000 Hz, per channel, using LabView (National Instruments, Austin, TX), then filtered in Matlab (The Mathworks, Natick, MA) using a phaseless 4th order Butterworth filter with a cut off frequency of 1650Hz. Angular acceleration measures were derived from the angular velocity data based on a 5-point least squares quartic equation.
About Lloyd Industries, Inc.
Lloyd Industries, Inc., located in San Antonio, Florida, is a research and development company focused on the biomechanics of brain injuries. The company was founded in 2004 by John D. Lloyd Bio, Ph.D., CPE, CBIS, Board Certified Ergonomist and Certified Brain Injury Specialist. He has also provided expert witness services nationwide for over 20 years in the fields of biomechanics, ergonomics and human factors, specializing in the biomechanics of brain injury, including sport and motorcycle helmet cases, slips and falls, motor vehicle accidents and pediatric head trauma. Lloyd Industries is open to licensing with manufacturers to bring this much-needed technology to market for the protection of sports participants and athletes of all ages. For additional information call 813-624-8986.
Dr. John Lloyd recently conducted a biomechanical study to evaluate motorcycle helmets in terms of their ability to provide protection against traumatic head and brain injuries. Motorcycle helmet testing proves inadequate protection against concussion and diffuse traumatic brain injuries associated.
Motorcycle accident victims account for more than 340,000 fatalities annually, with the United States ranking 8th highest worldwide in the number of motorcycle accident deaths. 75% of all fatal motorcycle accidents involve brain injury, with rotational forces acting on the brain the primary cause of mortality. Current motorcycle helmets are effective at reducing head injuries associated with blunt impact. However, the mechanism of diffuse traumatic brain injury is biomechanically very different.
Samples of 9 motorcycle helmet models, representing full-face, three-quarter and shorty designs were evaluated. Helmets, fitted to an instrumented Hybrid III head and neck, were dropped at 13 mph in accordance with DOT motorcycle helmet testing standards.
Results show that, on average, there is a 67% risk of concussion and a 10% probability of severe or fatal brain injury associated with a relatively minor 13mph helmeted head impact.
In conclusion, motorcycle helmets provide inadequate protection against concussion and diffuse traumatic brain injuries associated with even relatively moderate impact.
Motorcycle helmets were originally developed in the early 20th century and, like most helmets, are modeled after military helmets, the purpose of which is to protect against penetrating head injury. The modern motorcycle helmet, with a hard outer shell and rigid expanded polystyrene (EPS) liner was actually introduced over 60 years ago. The outer shell serves as a second skull, dispersing the impact force over a wider surface area, while the inner shell compresses in an attempt to reduce translational forces. A mechanism to mitigate tangential forces is absent. Since the liner fills the entire inner surface of the shell, tangential forces cannot be absorbed and are transmitted directly to the head and brain. Motorcycle helmet standards focus on reducing the effect of linear impact forces by dropping them from a given height onto an anvil and measuring the resultant peak linear acceleration.
Motorcycle Helmet Standards
In motorcycle helmet testing, the risk of impact loading injuries, such as skull fractures, can be determined by measuring linear accelerations experienced by a surrogate head form in response to impact. Whereas risk of impulse or inertial loading injuries, such as concussion, axonal injury and subdural hematoma can be quantified by measuring impact-related angular accelerations at the center of mass of a test head form.
Unfortunately, the evolution of motorcycle helmet design is not driven by advances in scientific knowledge, but rather by the need to meet applicable testing standards. In the United States, standards include the federal motor vehicle safety standard (FMVSS) #218, commonly known as the DOT motorcycle helmet testing standards, and Snell M2015, while ECE 22.05 and BSI 6658 were adopted in European countries. Test procedures involve dropping a helmeted head form onto various steel anvils at impact velocities ranging from only 5.0 to 7.75 m/s (11-17 mph). Pass/fail is based on the ability of the helmet to provide protection against forces associated with linear acceleration in response to impact.
Current motorcycle helmet testing standards do not incorporate measures of angular acceleration and therefore fail to assess whether helmets offer protection against catastrophic brain injuries. The omission of this critical measure is reflected epidemiologically in the disproportion of closed head injuries in fatal motorcycle accidents.
Human injury is complicated. If we lived our lives inside a protective bubble then, one day experienced an incident, it may be relatively simple to ascribe any injuries to the traumatic event. But that is typically not the case. As an aging nation, our bodies experience mechanical trauma every day – from work, sports, recreation and potential incidents. The question is whether forces and accelerations acting on the body as a result of a traumatic incident, such as an automobile collision, slip and fall, or recreational accident, were the direct and ultimate cause of injuries. Answering those questions is the unique role of forensic biomechanics and is typically beyond the expertise of most medical doctors.
Forensic biomechanics is the study of injury causation by measuring forces acting on and within the human body using methods of mechanics, to determine whether such forces exceed known thresholds of injury. As such, a biomechanist possesses expertise in the fields of both mechanics and human anatomy. Biomechanists and medical doctors serve complementary roles in the medico-legal system. Medical Doctors have specific knowledge to diagnose and treat a patient, however forensic biomechanics is not taught in medical school. Therefore, a biomechanist is required, based on their specialized education, training and experience, to serve as the necessary ‘bridge’ between medicine and engineering by calculating the forces acting on the body as a result of a claimed incident and thereby explaining the diagnosed injuries in terms of mechanical causation.
In a motor vehicle accident case, a biomechanist will assist the trier of fact by relating the impact forces and motions of the vehicles (automobiles, trucks, motorcycles, bicycles or pedestrians) to the resultant motion of occupants or other persons involved (kinematics). and forces they experience (kinetics) due to often multiple impacts within the vehicle interior or ground, then relate those forces to explain the mechanical causation of their medically diagnosed injuries.
Motorcycle, bicycle and pedestrian involved accidents can be substantially more complex, since the vehicles and operators tend to become separated and travel independently to their final rest positions. In Florida and many other states, motorcyclists have the right to choose whether or not to wear a helmet. Dr. Lloyd has conducted and published extensive research on the biomechanics of helmet protection, which shows that while helmets are effective at reducing the risk of penetrating head injury due to skull fracture, helmets do not offer adequate protection against traumatic brain injury, which can occur whether the rider is helmeted or not.
In a recent jury trial, Dr. Lloyd provided expert testimony in the fields of accident reconstruction, biomechanics and human factors on behalf of a plaintiff who suffered traumatic brain injury and a broken neck in a high-speed truck collision when a distracted driver drove through a stop sign. The jury awarded the plaintiff more than $14.5 million in damages.
Forensic biomechanics is also key to the analysis of cases involving slips, trips and falls, which are frequently claimed in all manners of environments, including workspaces, shopping arenas, restaurants, etcetera. Slips may occur whenever the coefficient of friction (CoF) between one’s footwear and flooring surface is too low, often due to the presence of a foreseeable foreign substance, such as a fluid. Whereas a trip may occur whenever the CoF between the footwear and flooring is too great, or unexpected, such as a transition between different surfaces. Unprotected falls can and do generate inordinate forces on the human body caused by acceleration due to gravity. For example, a simple fall from approximately 3 feet can generate an impact velocity of 10 miles per hour! But, more important is how quickly the human body comes to rest upon impact. It has been shown that a simple fall from only 12 inches onto a hard surface, such as concrete, can generate more than 1000 pounds of force on the human head, which is sufficient to cause fatal injury.
One recent slip and fall case in which Dr. Lloyd testified, involved a vascular surgeon, who went to sit down on a ‘budget’ stool to write his post-surgical notes. In that case, it was determined that the choice of casters on the wheeled stool were inappropriate for the environment, causing the stool to slip out from beneath the surgeon, who fell backwards, striking his head on the hard floor, resulting in traumatic brain injury and ongoing epileptic episodes. The surgeon, who suffered severe neurological deficits as a result of the incident was unable to return to work and also suffered many other lifelong effects. At trial the jury awarded the surgeon $10 million for injuries caused by the slip and fall.
In conclusion, a forensic biomechanical analysis may be pertinent to the success of a variety of cases, including: motor vehicle accidents (involving automobiles, trucks, motorcycles, bicycles and pedestrians), recreational accidents (including boating, jet skiing, ATVs, etc.), sports injuries / helmet protection as well as slips, trips and falls. The opinions formulated by Dr. Lloyd and other forensic biomechanists regarding the quantitative accelerations and forces necessary to result in injury are uniquely biomechanical opinions, and no other area of science or medicine is as appropriate to offer such opinions. Neither mechanical engineering nor physics include the prerequisite background concerning human body tissue properties and human anatomy. Similarly, medical training does not provide the necessary understanding of biomechanical principles to identify qualitative relationships between physical trauma and human tissue injury. Thus, a forensic biomechanist serves the legal system by quantifying the forces associated with an incident and comparing those forces against scientifically accepted thresholds of injury thereby explaining the medical diagnosis.
Judge Healey of the State of Florida First District Court of Appeals (Case No. 1D11-4210) recently upheld the importance of forensic biomechanics testimony in his ruling, which stated that “a biomechanics expert is qualified to offer an opinion as to causation if the mechanism of injury falls within the field of biomechanics” and as such is “relevant to establishing a reasonable hypothesis … that the victim’s injuries were consistent with … trauma”.
Ultimately, the success of any expert lies in their ability to convey often complex matters to a jury. Based on over 20 years of experience as an expert, during which time Dr. Lloyd has provided testimony at trial or in deposition more than 80 occasions, he has become highly proficient in using methods that express complex matters in simplistic terms for the purpose of educating the jury as to the facts of a case.
The following is a case study in which biomechanics expert, Dr. John Lloyd, evaluated the risk of concussion and brain injury associated with headrest impact in rear end crashes.
Headrest Impact Test Apparatus:
In accordance with prior published test methods[1],[2],[3], a test apparatus was constructed to evaluate the biomechanical protection afforded by an exemplar automobile headrest against head and brain injuries during occipital head impacts in a simulated rear-end motor vehicle collision.
The apparatus involves a pendulum arm, attached by bearing housings to a weighted base. The upper body, including neck and head of a 50th percentile Hybrid III crash test dummy was mounted to the pendulum arm. Data acquisition was initiated by triggering an electromechanical release mechanism, allowing the mannequin to fall, under acceleration due to gravity, until the crash test dummy impacted the headrest and backrest (Figure 1).
Figure 1: Test apparatus
The fundamental elements and principles of this testing have been utilized in other laboratories. By utilizing a Hybrid III neck, the head impact tests are more realistic, causing head rotation at the axis between the head and neck, which produces measures of head and brain angular kinematics. The methods presented herein are based upon standardized test methodologies and published research.
Instrumentation
Four PCB Piezotronics tri-axial accelerometers (model # 356A01) were mounted in an X,Y,Z array at the center of mass of the Hybrid III headform, along with a tri-axial angular rate sensor produced by Diversified Technical Systems (composite Figure 2).
Figure 2: Sensor installation in Hybrid III headform
Sensor Calibration:
All sensors were calibrated by the manufacturer. Verification of calibration of the linear accelerometers was performed prior to testing using a calibration shaker. Results indicate that the sensors were operating in the specified frequency range and output (Figure 3).
Figure 3: Pre-test verification of linear accelerometer sensors
For the angular rate sensor, a simple validation method was devised in which the sensor was affixed to a digital goniometer that was rotated through a 90-degree angle. Using LabVIEW software, the integral of angular rate was computed, reflecting concurrence with the digital goniometer for all three planes of motion (Figure 4).
Figure 4: Pre-test validation of angular rate sensor calibration
Headrest Impact Testing:
The mannequin head was raised from the headrest in 2-inch increments from 2 inches to 30 inches, generating head impact speeds from 1 to 25 miles per hour. Two headrest positions were evaluated, along with two different Hybrid III necks representative of a stiff and relaxed neck (Figure 5), for a total of sixty tests.
Figure 5: Test apparatus with Hybrid III loose neck and headrest in lower position
Data Acquisition and Analysis:
Data from the analog sensors were acquired in accordance with SAE J211 [4], using a National Instruments compact DAQ data acquisition system and LabVIEW software (National Instruments, Austin, TX). The raw data was then filtered in MATLAB (The MathWorks, Natick, MA) using a phaseless eighth-order Butterworth filter with cutoff frequencies of 1650 Hz and 300Hz for the linear accelerometers and angular rate sensors, respectively.
Angular acceleration values for sagittal, coronal and axial planes were computed from the angular velocity data using the 5-point central difference by least squares method (Equation 1):
Equation 1: Five-point central difference by least squares method
Angular acceleration vales were also derived from the array of linear accelerometers, by the mathematical method documented by Padgaonkar et al [5].
Linear velocity was calculated by integrating linear acceleration. Mathematical methods were performed using Matlab to compute characteristic values from variables of interest. Figure 6, below illustrates peak linear acceleration and angular velocity associated with a 6.8 mph occipital head impact against a headrest.
Figure 6: Linear acceleration and angular velocity associated with headrest impact
It is noted that the major component of linear acceleration was in the X-axis (anterior-posterior), while the major component of angular velocity was in the sagittal plane, as expected.
Linear acceleration values were used to calculate Maximum Pressure (Equation 2), Gadd Severity Index (GSI) (Equation 3), and Head Injury Criterion (HIC15) (Equation 4).
Equation 2: Maximum Pressure
Equation 3: Gadd Severity Index
The Head Injury Criterion (HIC) is an empirical measure of impact severity describing the relationship between the linear acceleration magnitude, duration of impact and the risk of head trauma (Equation 4).
Equation 4: Head Injury Criterion
where a is resultant head acceleration, t2-t1 < 15 msec
With reference to the Figure 7, below, the HIC value is used to predict the risk of head trauma: Minor –skull trauma without loss of consciousness; nose fracture; superficial injuries Moderate – skull trauma with or without dislocated skull fracture and brief loss of consciousness. Fracture of facial bones without dislocation; deep wound(s) Critical – Cerebral contusion, loss of consciousness for more than 12 hours with intracranial hemorrhaging and other neurological signs; recovery uncertain.
Figure 7: Probability of specific head trauma level based on HIC value
Peak angular velocity was determined as the maximum angular velocity related to peak linear acceleration impact time. Angular velocity values were used to derive Maximum Principal Strain (MPS) (Equation 5), Cumulative Strain Damage Measure (CSDM) (Equation 6), and Brain Rotational Injury Criterion (BrIC) (Equation 7).
Equation 5: Maximum Principal Strain
Equation 6: Cumulative Strain Damage Measure
An analysis method validated by Takhounts [6] establishes physical injury criteria for various types of traumatic brain injury and uses Anthropomorphic Test Device (ATD) data to establish a kinematically based brain injury criterion (BrIC) for use with ATD impact testing. This method was utilized to express risk of diffuse brain injury according to the revised AIS scale [7] in terms of peak angular head kinematics, where:
Equation 7: Brain Rotational Injury Criterion
Headrest Impact Results:
A summary of key results is presented in Table a-d, below. The driver was aware of the pending impact, as he depressed the accelerator in an attempt to avoid the collision in the moments prior to the crash. In rear end collision tests involving human subjects, volunteers instinctively tensed their neck muscles as a protective response. Given that the driver anticipated the crash his neck muscles were likewise expectedly tense as an instinctive protective response. Therefore, the results most consistent with the subject case are presented in Tables a and b. Rows highlighted in green are consistent with change in velocity experienced by the driver during the subject crash.
Table a: Summary of test results – Neck – Stiff; Headrest – lower position
Table b: Summary of test results – Neck – Stiff; Headrest – upper positio
Table c: Summary of test results – Neck – Loose; Headrest – lower position
Table d: Summary of test results – Neck – Loose; Headrest – upper position
Skull Fracture
With reference to Ono 8, none of the impact tests approached the occipital skull fracture threshold of 140 g for impacts lasting longer than 7 milliseconds. Therefore, vehicle headrests provide excellent protection against acute skull fractures at impact speeds below 25 mph.
Traumatic Head Injury
With reference to Figure 7 and Tables a-d, maximum recorded HIC values were consistent with a 5 percent or less risk of moderate traumatic head injury. Whereas, the HIC value computed at impact speeds similar to the crash was only 3.4, at which the risk of minor or moderate traumatic head injury is negligible.
Mild Concussion
With reference to Figure 8 below, the risk of an occupant sustaining a mild concussion in a rear-end collision producing a change in velocity of 6.25 mph (range 5.4 to 7.2 mph) can be determined based on the following calculation: Risk AIS-1 = 31.744*ln(x) + 6.1748 (R2=0.67). The risk of and AIS-1 mild concussion, without post-concussion syndrome, in such an impact is 64.3% (range 59.7 to 68.8%).
Figure 8: Risk of mild concussion (AIS-1) associated with headrest impact
Severe Concussion
With reference to Figure 9, below, the risk of an occupant sustaining a severe concussion in a rear-end collision producing a change in velocity of 6.25 mph (range 5.4 to 7.2 mph) can be determined based on the following calculation: Risk AIS-2 = 0.198e0.234x (R2=0.85). The risk of severe concussion in such an impact is 0.85% (range 0.70 to 1.07%).
Figure 9: Risk of severe concussion (AIS-2) associated with headrest impact
Traumatic Axonal Injury:
Figure 10, below, is adapted from Margulies et al. 20 in which thresholds for axonal injury were developed and published based on mathematical modeling, animal testing and physical experiments. Results from occipital head impact against an exemplar headrest at a speed of 6.2 miles per hour are represented, indicating that rotational head and brain kinematics associated with such impact are well below scientifically-accepted thresholds for traumatic axonal injury.
Figure 10: Scientific Thresholds for Axonal Injury
Figure 11, below was generated from data presented in Tables a through d, to present the risk of traumatic axonal injury associated with head impact against an headrest.
Figure 11: Risk of traumatic axonal injury (AIS-4) associated with headrest impact
Results show that the risk of an occupant sustaining traumatic axonal injury in a rear-end collision producing a change in velocity of 6.25 mph (range 5.4 to 7.2 mph) can be determined based on the following calculation: Risk AIS-4 = 0.0271e0.2391x (R2=0.85). The risk of traumatic axonal injury in an impact of the magnitude experienced by the driver is 0.12% (range 0.10 to 0.15%).
Conclusions
Biomechanical testing of head and brain injury risk associated with occipital head impact against a headrest, in accordance with published methods, shows a significant risk (59.7 to 68.8%) of AIS-1 mild concussion, without post-concussion syndrome, in a 6.2 mph rear-end collision. However, the risk of an AIS-2 severe concussion in such an impact decreases to 0.70 to 1.07%, and the risk of traumatic axonal injury is only 0.10 to 0.15%. Moreover, the mechanical traumatic axonal injury is not consistent with a sagittal plane impact.
References
[1] Caccese V, Lloyd J, Ferguson J (2014) An Impact Test Apparatus for Protective Head Wear Testing Using a Hybrid III Head-Neck Assembly. Experimental Techniques.
[2] Lloyd J & Conidi F. (2015). Brain Injury in Sports. Journal of Neurosurgery. October.
[3] Lloyd J. (2017). Biomechanical Evaluation of Motorcycle Helmets: Protection Against Head and Brain Injuries.Journal of Forensic Biomechanics.
[4] SAE (2014) J211/1. Instrumentation for Impact Test – Part 1 – Electronic Instrumentation. Society of Automotive Engineers International, Surface Vehicle Recommended Practice, Warrendale, PA.
[5] Padgaonkar AJ, Krieger KW and King AI. Measurement of Angular Acceleration of a Rigid Body using Linear Accelerometers. J Applied Mechanics. Sept 1975.
[6] Takhounts EG, Craig MJ, Moorhouse K, McFadden J (2013) Development of Brain Injury Criteria (BrIC). Stapp Car Crash Journal 57: 243-266.
[7] Abbreviated Injury Scale (2008) Association for the Advancement of Automotive Medicine, Des Plaines, IL.
* peer-reviewed and published in the Journal of Forensic Biomechanics, October 2017 ** DOWNLOAD PDF FILE
Abstract
Motorcycle accident victims worldwide account for more than 340,000 fatalities annually, with the Unites States ranking 8th highest in number of motorcycle accident deaths, largely due to non-mandatory motorcycle helmet requirements for adults in a number of States. Seventy-five percent of all fatal motorcycle accidents involve head and brain injury, with rotational forces acting on the brain the primary cause of mortality. Current motorcycle helmets are reasonably effective at reducing head injuries associated with blunt impact. However, the mechanism of traumatic brain injury is biomechanically very different from that associated with focal head injury. This study was conducted to evaluate the effectiveness of current motorcycle helmets at reducing the risk of traumatic brain injuries.
Ten motorcycle helmet designs, including full-face, three-quarter and half-helmets were evaluated at an average impact velocity of 8.3 ms-1 (18.5 mph) using a validated test apparatus outfitted with a crash test dummy head and neck. Sensors at the center of mass of the headform enabled high-speed data acquisition of linear and angular head kinematics associated with impact.
Results indicate that none of the standard helmet models tested provide adequate protection against concussion and severe traumatic brain injuries at moderate impact speeds. Only one of the standard motorcycle helmet models tested provided adequate protection against skull fracture.
A new motorcycle helmet prototype, incorporating a liner constructed from a composite matrix of rate-dependent materials was tested, with comparison to standard motorcycle helmet designs, with very promising results. Knowledge learned from this study will facilitate the development of a new generation of advanced motorcycle helmets that offer improved protection against both head and brain injuries.
In developing countries motorcycles are required for utilitarian purposes due to lower prices and greater fuel economy, whereas in the developed world they are considered a luxury and used mostly for recreation. In 2016 there were more than 134 million motorcycles worldwide [1], 8.4 million of which were registered in the United States, representing 3.2% of all US registered vehicles. California, Florida and Texas were the leading states in terms of the motorcycle popularity; collectively representing 22% of all US registered motorcycles [2]. In 2011, U.S. motorcyclists travelled a total of 18.5 billion miles, which, while only 0.6% of total vehicle miles travelled, accounted for 14.6% (4,612) of U.S. traffic fatalities that year. Worldwide there are more than 340,000 motorcyclist fatalities annually, which equates to more than 28% of all road accident deaths [3]. According to the U.S. National Highway Traffic Safety Administration (NHTSA) and other reports, when compared per vehicle mile traveled with automobiles, due to their vulnerability, motorcyclists’ risk of a fatal crash is 30-35 times greater than that of a car occupant [4][5][6][7].
Two fundamental epidemiologic studies into the causation of motorcycle accidents have been conducted: the Hurt study in North America and the MAIDS study in Europe. According to the Hurt Report [8], 75 percent of collisions were found to involve a motorcycle and a passenger vehicle, while the remaining 25% were single vehicle accidents. The cause of motorcycle versus passenger vehicle collisions in 66% of accidents involves violation of the rider’s right of way due to the failure of motorists to detect and recognize motorcycles in traffic. Findings further indicate that severity of injury to the rider increases with alcohol consumption, motorcycle size and speed.
The most recent epidemiologic study to investigate motorcycle accident exposure data was conducted between 1999-2001 by a partnership of five European countries [9]. Findings show that passenger cars were again the most frequent collision partner (60%), where more than two-thirds of drivers reported that they did not see the motorcycle and more than half of all accidents involving motorcycles occurred at an intersection.
The COST report, which is an extension of the MAIDS study, documents that three-quarters (75%) of all motorcyclist deaths are a result of injury to the head and brain [10]. Linear forces were the major factor in 31% of fatal head injuries, while rotational forces were found to be the primary cause in over 60% of cases. While the helmet is considered the most effective means of rider protection [11], recent studies indicate that motorcycle helmets are only 37-42% successful in preventing fatal injury [12],[13]. By reducing peak linear forces acting on the head it was commonly believed that the risk of diffuse brain injuries, including concussion, subdural hematoma and diffuse axonal injury would also be prevented [8]. However, the biomechanical mechanisms of head and brain injuries are unique. New research shows that these mechanisms are poorly correlated [14].
Motorcycle Helmet Standards
Like most helmets, motorcycle helmets are modeled after ancient military helmets, the purpose of which is to provide protection against penetrating head injury, such as skull fracture. Whereas, all impacts have both linear and oblique components, which produce translational and tangential forces, respectively. The modern motorcycle helmet was introduced over 60 years ago [15]. Its outer shell serves as a second skull, diffusing impact forces over a larger surface area, while the inner liner compresses to minimize translational forces. However, a mechanism to mitigate tangential forces is absent. Since the liner fills the entire inner surface of the shell and is immobile, rotational inertia induced tangential forces are transmitted directly to the brain.
The likelihood of a helmeted motorcyclist sustaining impact loading injuries, such as skull fractures, can be determined by quantifying the magnitude of peak linear acceleration experienced by a test headform in response to impact. Whereas the risk of a rider suffering inertial or impulse loading injuries, such as concussion, axonal injury and intracranial hematoma can be computed based on impact-related angular kinematics at the headform center of mass [16],[17].
Unfortunately motorcycle helmet protection is not driven, for the most part, by advances in scientific knowledge, but by the need to meet applicable testing standards [18],[19]. In the United States, the governing specification is the federal motor vehicle safety standard (FMVSS) #218 [20]; the Snell Memorial Foundation also offers a voluntary standard M2015, which is a little more stringent [21]. Whereas BSI 6658 [22] and ECE 22.05 [23] have been adopted in European countries and AS/NZS 1698 accepted in Australasian countries [24]. Test protocols involve the guided fall of a helmeted headform onto steel anvils of various designs at impact velocities ranging from only 5.2 to 7.5 m/s (11-17 mph). The pass/fail criterion is based only on the helmet’s effectiveness in reducing peak linear acceleration, and thereby translational forces, in response to impact.
Impact-related angular head kinematics are not quantified under current motorcycle helmet standards, which therefore fail to assess whether helmets offer any protection against traumatic brain injuries. The omission of this critical measure of helmet performance is reflected epidemiologically in the disproportion of closed head and brain injuries in fatal motorcycle accidents [9,10].
Biomechanics of Head and Brain Injury
The two mechanisms associated with traumatic head and brain injury are impact loading and impulse loading, both of which are present in all impact events. Impact loading involves a blow directed through the center of mass of the head, resulting in translation of the head and brain. When thresholds of injury are exceeded, skull fractures [25], lacerations and contusions (bruising) to the head and underlying brain tissue may result [26]. Whereas, impulse or inertial loading is produced when an oblique impact, common to motorcycle crashes, creates tangential forces, causing head rotation. Since the brain is not rigidly attached to the inside of the skull, rotational inertia of the brain produces a mechanical strain on cerebral blood vessels, nerve fibers and brain tissue. When thresholds of injury are exceeded, nerve fibers in the brain may be damaged, producing concussion [27] and diffuse axonal injury (DAI) [28]. Blood vessels may also rupture, causing subdural hemorrhages (SDH) [29], the high mortality rate of which has motivated numerous studies of bridging vein failure properties [30],[31],[32],[33],[34],[35]. Subdural hematoma and traumatic axonal injury are frequently identified as the cause of serious injury or fatality in motorcycle accidents.
Holbourn [[36]] was the first to identify angular / rotational acceleration as the principal mechanism in brain injury. Gennarelli, Ommaya and Thibault further investigated the importance of rotational (angular) acceleration in brain injury causation, based on studies involving live primates and physical models, [28,29,[37],[38],[39], concluding that angular acceleration is far more critical than linear acceleration to the causality of traumatic brain injuries. They further isolated and investigated the unique effects of translational (linear) and inertial (angular) loading on the heads of primates [28], confirming that pure translation produces focal injuries, such as contusions and skull fractures, while rotationally induced inertial loading causes diffuse effects, including concussion and subdural hematoma. Closed head and brain injury, found in more than 60% of motorcycle accident fatalities, is due to inadequate helmet protection against impact-related angular head kinematics [10].
Skull fracture:
Ono [25] published thresholds for human skull fracture based on cadaver experiments. Twenty-five human cadaver skulls were exposed to frontal, occipital and lateral impacts. Each skull was covered with the rubber skin of a Hybrid II mannequin and filled with gelatin to accurately represent head mass. A series of 42 frontal, 36 occipital and 58 temporal blows were delivered to the suspended heads, during which linear accelerations were measured. A skull fracture threshold of 250 g for 3-millisecond impulse duration was established for frontal and occipital impacts, decreasing to 140 g for 7-millisecond impulse duration. Whereas the skull fracture threshold for lateral impacts is reported as 120 g over 3-millisecond duration, decreasing to 90 g over 7 milliseconds. Results indicate that skull fracture threshold is inversely related to impulse duration.
Concussion:
Several studies have attempted to establish biomechanical thresholds for concussion. Pellman et al. analyzed a series of video-recorded concussive impacts during NFL football games, reporting that concussive injury is possible at 45 g / 3500 rad/s2, while 5500 rad/s2 represents a 50% risk of concussive trauma [40]. Rowson and Duma, also using head injuries in America football as their model, conducted extensive laboratory and field-based biomechanical evaluations [41],[42],[43],[44]. Based on data from 62,974 sub-concussive impacts and 37 diagnosed concussions recorded using the Simbex, Inc. (Lebanon, NH) Head Impact Telemetry System (HITS), the investigators propose a concussion threshold of 104 ± 30 g and 4726 ± 1931 rad/s2.
Subdural Hematoma:
According to Gennarelli, the most common form of acute subdural hematoma (ASDH) is caused by shearing of veins that bridge the subdural space [29]. The severity of injury associated with bridging vein rupture has led to numerous studies of their mechanical properties (Lowenhielm [30-31,32], Lee and Haut [33], Meaney [34], and Depreitere [35]).
Lowenhielm tested 22 human parasagittal bridging vein samples from 11 decedents between the ages of 13 and 87 years without history of brain injury [30,31]. He hypothesized that blunt trauma to the head causes the brain to be displaced with respect to the dura, thereby stretching bridging veins and surrounding connective tissue. Based on his laboratory experiments, Lowenhielm found that maximal shear stresses occur about 7 milliseconds after impact, coinciding with bridging vein disruption. He concluded that bridging vein rupture may occur if peak angular acceleration exceeds 4500 rad/s2.
Depreitere subjected ten unembalmed human cadavers to 18 occipital impacts producing head rotation of varying magnitude and impulse duration in the sagittal plane [35]. Bridging vein ruptures, detected by autopsy, were produced in six impact tests. Findings suggest a mean tolerance level of approximately 6,000 rad/s2 for 10-millisecond impulse duration, which seems to decrease for longer impulse durations, however the confidence interval is rather broad due to the limited data set. Data from the research by Depreitere and Lowenhielm is presented in Figure 1.
Figure 1: Bridging vein failure as a function of impulse duration and peak angular acceleration (with line of best fit and 75% confidence intervals).
Helmets decrease peak translational force by extending the impulse duration. In the case of motorcycle helmets, typical impulse duration is approximately 12 milliseconds. With reference to Figure 1, above, this suggests that bridging vein rupture may result with peak angular accelerations in the order of 5,000 rad/s2, but may be as low as 3,000 rad/s2 after adjusting for standard error of the mean in this limited dataset.
While previous studies have investigated motorcycle impacts into vehicles and fixed barriers, the underlying motivation of such studies was to determine crush characteristics of the vehicles for accident reconstruction purposes [45]. Other studies have evaluated peak linear accelerations of the head, chest and pelvis of motorcyclists in collisions [46]. However, rotational forces associated with impact-related peak angular accelerations have not been determined even though it is well known that rotational mechanisms are the primary cause of closed head injuries [28,29,36-37,38,39] in helmeted motorcyclist accidents [10]. Measurement of impact-related head angular / rotational acceleration is critical to the development and evaluation of motorcycle helmets to provide effective protection against traumatic brain injuries associated with a range of typical motorcycle crash-related head impact speeds. To that end, this paper offers an objective determination of the performance of a variety of motorcycle helmets in terms of their ability to protect against both head and traumatic brain injuries associated with impact velocities reflective of typical head impact velocities in motorcycle accidents.
Methods
The standard test apparatus for impact testing of protective headwear was modified to enable measurement of both linear and angular headform kinematics [16]. This validated apparatus is comprised of parallel vertical braided stainless steel wires that guide the fall of a 50th percentile Hybrid III head and neck assembly (HumaneticsATD, Plymouth, MI) mounted to an aluminum flyarm. The anvil onto which the headform impacts consists of a 50 mm thick steel base plate, with a 100 mm thick concrete overlay, consistent with the coefficient of friction for typical roadway surfaces. Figure 2 illustrates this setup.
Figure 2: Modified Head drop system with Hybrid III head / neck
According to Mellor et al. [47] apparatus for the evaluation of protective headgear in which the headform is rigidly affixed to the carriage (flyarm) reduces the dissipation of energy by excessive rotation of the helmeted headform and sliding of the helmet on the anvil, thereby inflating peak linear acceleration measures. Examples in which the headform is rigidly affixed to the flyarm include the FMVSS218 test apparatus [20]. Whereas in Snell M2015 [21], BS 6658 [22] and AS/NZS 1698 [24] specifications the headform is attached to the flyarm by means of a hinge joint, which allows headform rotation in the sagittal plane as well as vertical translation, but prevents motion in the coronal and axial planes. The ECE 22:05 test method [23] utilizes a ball joint between the flyarm and headform, thereby permitting unrestricted head rotation in all three planes. Similar to the ECE test method, utilization of the Hybrid III neck permits headform rotation in sagittal, coronal and axial planes, but limits the rate of motion in a manner more consistent with the human musculoskeletal system [48]. Moreover, orientation of the Hybrid III neck was maintained relative to the flyarm, irrespective of headform orientation, thereby standardizing response of the neck form.
Instrumentation: A triaxial block, installed at the center of mass of the Hybrid III headform (HumaneticsATD, Plymouth, MI) housed a triaxial accelerometer from PCB Piezotronics (Depew, NY) and three DTS-ARS Pro angular rate sensors (Diversified Technical Systems, Seal Beach, CA). Data from the sensors were acquired using compact DAQ hardware from National Instruments (Austin, TX).
While all sensors had been calibrated by the respective manufacturers, verification tests were performed to validate linear and angular sensor calibration data. Calibration of the tri-axial linear accelerometer was validated using a portable handheld shaker and found to be within specification for all three axes of measurement. For the angular rate sensor a simple validation method was devised in which the sensor was affixed to a digital goniometer, which was moved through a set angle (Figure 3). Using LabView, the integral of angular rate was computed, reflecting concurrence with the digital goniometer for all three planes of motion.
Figure 3: Validation of Angular Rate Sensor Calibration
Ten motorcycle helmet models were selected for evaluation, based on popularity among motorcyclists, including representative models of full-coverage, three-quarter and half-helmet (shorty) styles, as shown in Figure 4, below. All models displayed the DOT certification sticker, indicating that their protective performance met the FMVSS218 motorcycle helmet testing standard [20]. Helmet sizes were chosen based on best fit for the Hybrid III headform, which has a 58cm head circumference, representative of a 50th percentile US adult male.
Figure 4: Motorcycle Helmet Models Evaluated
In addition, a new prototype motorcycle helmet (Figure 5) was tested for comparison against the ten standard DOT motorcycle helmets. The prototype helmet was a three-quarter standard shell with liner constructed from a composite of rate-dependent materials arranged in a patent-pending matrix [49].
Figure 5: Motorcycle Helmet Prototype
Five samples of each motorcycle helmet model were purchased in new condition. Each helmet was impacted one time in the frontal and/or occipital region at an impact velocity of approximately 8.3 meters per second (18.5 mph), which was verified computationally. Repeatability of the tests was confirmed at the start and end of data collection by dropping the Hybrid III headform from a height of 2.0 m onto a Modular Elastomer Programmer (MEP) pad of 25 mm thickness and durometer 60A. Standard Error of the Mean of 0.061 was computed based on peak angular accelerations for pre and post MEP pad drop tests.
Analysis: Analog sensor data were acquired at 20 kHz per channel, in accordance with SAE J211 [50], using LabView (National Instruments, Austin, TX). The raw data was then filtered in MATLAB (The MathWorks, Natick, MA) using a phaseless eighth-order Butterworth filter with cutoff frequencies of 1650 Hz and 300Hz for the linear accelerometers and angular rate sensors, respectively. Angular acceleration measures were computed from the angular velocity data using 5-point least-squares quartic equations. Impulse duration was determined based on the linear acceleration signal, where impulse start point is the time at which the magnitude of linear acceleration exceeds 3 g and impulse end point is the time at which the major component of linear acceleration crosses the y-axis (Figure 6). The gradient from impulse start point to peak was computed, as was the area under the acceleration magnitude curve from start to end points. Variables for the angular acceleration signal were similarly computed.
Figure 6: Impulse duration based on linear acceleration data
An analysis method validated by Takhounts [51] establishes physical (strain and stress based) injury criteria for various types of brain injury based on scaled animal injury data and uses Anthropomorphic Test Device (ATD) test data to establish a kinematically based brain injury criterion (BrIC) for use with ATD impact testing. This method was utilized to express risk of brain injury according to the recently revised AIS scale [52] in terms of peak angular head kinematics, where:
The probability of brain injury for AIS 1-5 was thus computed as a function of BrIC:
Additionally, mechanical head and brain injury parameters of maximum pressure (in kPa), maximum principal strain (MPS) and cumulative strain damage measure (CDSM) were computed for each helmet impact test:
Results
The following table presents a summary of results for each of helmet models evaluated:
Table 1: Summary of Results
* The best performing helmet for each variable is highlighted in green
* The worst performing helmet for each variable is highlighted in red
Motorcycle Helmet Protection against Skull Fracture:
Figure 7, below, presents peak linear acceleration values, averaged across 5 samples of each of the 10 motorcycle helmet models tested, along with results for the prototype, against pass/fail thresholds for current motorcycle helmet testing standards (DOT, Snell, BS and ECE) as well as frontal-occipital and lateral skull fracture thresholds, per Ono [25].
Figure 7: Risk of Skull Fracture Associated with Motorcycle Helmet Impacts
Results show that while all of the motorcycle helmet models evaluated satisfy at least the DOT standard, only the Scorpion T510 full-face helmet offers sufficient protection against fronto-occipital and lateral impacts at the moderate impact velocities at which the helmets were tested.
Motorcycle Helmet Protection against Concussion:
Figure 8 presents peak angular acceleration results for 8.3 m/s impacts onto a concrete anvil, averaged across 5 samples of each helmet model. The red horizontal line on figure 8 indicates the 50% threshold for concussive trauma, as defined by Pellman et al [40].
Figure 8: Risk of Concussion Associated with Motorcycle Helmet Impacts
Results show that while a DOT approved motorcycle helmet may reduce peak angular acceleration associated with a helmeted head impact, the level of protection is not sufficient to prevent concussive injury in a typical motorcycle accident. Only the prototype motorcycle helmet, incorporating a liner constructed from a composite of rate-dependent materials arranged in a patent-pending matrix [49], offered adequate protection against concussive events.
Motorcycle Helmet Protection against Subdural Hematoma:
Figure 9, below, presents peak angular acceleration as a function of impulse duration, averaged across 5 samples of each of the 10 motorcycle helmet models tested, along with results for the prototype helmet. The threshold for bridging vein failure and resultant subdural hematoma is represented by the black line of best fit. Upper and lower boundary limits of this threshold are indicated in red, which represents a 75% likelihood that a subdural hematoma may occur for peak angular accelerations above the lower red line.
Figure 9: Risk of Subdural Hematoma Associated with Motorcycle Helmet Impacts
Most of the helmets tested, with exception of the prototype, fall above the lower threshold line suggesting the likelihood of catastrophic brain injury associated with a moderate helmeted impact. In fact, all but one of the five half-helmet models tested produced results above the mean threshold for subdural hematoma, indicating a higher likelihood of severe (AIS 4) or critical (AIS 5) brain injury. Overall, it appears that full-face helmets generally outperform half helmets in reducing the risk of subdural hematoma. Interestingly, an unhelmeted individual can seemingly withstand substantially greater peak angular accelerations and consequently experiences a lower risk of catastrophic brain trauma than a helmeted individual.
Correlation Analyses:
Pearson’s correlations were computed between each of the variables. Trends were suggested if computed R2 values were greater than 0.70, while strong correlations are indicated if R2 exceeded 0.80. Across all measures, the three most important variables, in rank order, for determining risk of head and brain injury are peak angular acceleration, angular acceleration gradient, and area under the angular acceleration curve between impulse start to end. The following interesting results were observed:
A negative trend exists between helmet mass and both linear acceleration (-0.70) and angular acceleration (-0.72). That is, both peak linear acceleration and peak angular acceleration seem to decrease as helmet mass increases.
There is neither a trend nor strong correlation between linear velocity and any of the variables investigated. This finding suggests that risk of head and brain injury is not related to impact speed.
A strong negative correlation exists between peak linear acceleration and impulse duration (-0.92). That is, impulse duration increases as peak linear acceleration decreases.
A trend, but not strong correlation was found between peak linear acceleration and peak angular acceleration, indicating that reducing impact-related peak linear acceleration may not necessarily mitigate peak angular acceleration.
Peak angular acceleration is strongly correlated with rotational injury criterion (RIC36) (0.95), Brain rotational Injury Criterion (BrIC) (0.93), probability of brain injury AIS 2 through 5 (μ=0.91), angular acceleration gradient (0.98), and area under the angular acceleration curve (0.96). A strong negative correlation is identified between peak angular acceleration and cumulative strain damage measure (CSDM) (-0.94) and maximum principal strain (MPS) (-0.94). A positive trend is also noted between peak angular acceleration and maximum pressure (0.77), Gadd Severity Index (GSI) (0.74) and linear acceleration gradient (0.76).
Discussion
As presented, the mechanisms associated with causation of focal head injuries and diffuse brain injuries are very different. Helmets were originally intended and continue to be designed to reduce the risk of potentially fatal head injuries caused by skull fracture fragments penetrating the brain. While skull fractures have been almost entirely eliminated in activities such as American Football, the higher impact speeds associated with motorcycle collisions continue to result in life-threatening cranial fractures, even in areas covered by the helmet. Thus, minimizing peak linear accelerations remains an important function of any motorcycle helmet. Therefore, to minimize the risk of skull fractures associated with helmeted motorcycle collision, based on research by Ono [25], a threshold of 140 g for peak linear acceleration to the frontal and occipital areas of the head and 90 g for peak linear acceleration for lateral impacts is suggested as a suitable performance criteria.
However, as with most helmets, motorcycle helmets perform inadequately in terms of mitigating the forces responsible for causing traumatic brain injury. Though a trend may exist between peak linear acceleration and peak angular acceleration, a strong correlation is absent, consistent with prior work in this area [14]. Hence, reduced peak linear acceleration through improved helmet design may not reduce the risk of traumatic brain injury. Indeed, as results herein show, an unhelmeted individual may be at a lesser risk of subdural hematoma during a moderate speed impact than one who is wearing a DOT approved motorcycle helmet.
Motorcycle Helmet Biomechanics
To minimize the risk of traumatic brain injury, spanning from mild concussion (AIS2) through severe brain injury (AIS5), it is necessary to reduce impact-related peak angular velocities in the sagittal, coronal and axial planes. Furthermore, since risk of subdural hematoma is defined based on peak angular acceleration and impulse duration, reducing peak angular velocities while also managing impulse duration will also lend to risk reduction of such severe or critical traumatic brain injuries. Therefore, to minimize the risk of concussion and subdural hematoma in helmeted motorcycle collisions, it is suggested that performance criteria based on peak angular velocity and acceleration not exceed 15.0 rad/s and 3,000 rad/s2, respectively, as previously proposed for American Football helmets [17].
Figure 10, below, was prepared to illustrate the relative effectiveness of the ten motorcycle helmet models tested and prototype in terms of protection against skull fracture, concussion and subdural hematoma, based on the above suggested performance criteria. Results indicate that only the prototype provides adequate protection against both traumatic head and brain injuries.
Figure 10: Motorcycle Helmet Effectiveness in Protecting Against Skull Fracture, Concussion and Subdural Hematoma
Based on the overall performance in terms of protection against skull fracture, concussion and subdural hematoma, and assuming equal weighting of these criteria for visualization purposes, the helmet models are presented in rank order in Figure 11.
Figure 11: Motorcycle Helmet Effectiveness (presented in rank order from left to right)
A strong negative correlation has been shown between helmet mass and both peak linear and angular accelerations. This finding suggests that ‘novelty’ motorcycle helmets (i.e. those not meeting FMVSS218 or other motorcycle helmet standards), which are often of lighter weight than DOT-approved helmets, will likely perform poorly in terms of preventing both head and brain injuries.
The new motorcycle helmet prototype evaluated within the scope of this study demonstrated exceptional potential to minimize the risk of traumatic brain injury, from mild concussion through severe brain injury, for a helmeted motorcyclist involved in a collision of moderate head impact speed.
Conclusions
The purpose of a motorcycle helmet is to reduce blunt force trauma to the head, thereby decreasing the risk of lacerations, contusions and skull fractures,. Whereas brain injuries may be produced when the brain lags behind sudden head motion thereby causing brain tissue, nerves and blood vessels to stretch and tear. The type of brain injury sustained is dependent on the magnitude and the time (pulse) duration over which mechanical stresses and strains act on the brain.
Motorcycle helmet test standards focus on reducing forces associated with linear acceleration by dropping helmeted headforms onto an anvil from a stated height and measuring the resultant peak linear acceleration. In general, the helmet design is considered acceptable if the magnitude of peak linear acceleration is less than an established threshold. Thus, helmets can and do prevent fatalities associated with penetrating head trauma. However, it may be argued that protection against brain injury is of paramount importance. After all, cuts, bruises and even bone fractures will heal, but brain injuries, if not fatal, often have life long neurologically devastating effects.
Current helmet testing standards do not require performance measures in terms of angular head kinematics and therefore fail to address whether motorcycle helmets provide the necessary protection against traumatic brain injuries. Research presented herein shows that it is possible to sustain catastrophic brain injuries, even while wearing a motorcycle helmet certified according to present testing standards.
Future generations of motorcycle helmets ought to be evaluated at higher impact velocities that are more indicative of head impact velocities in typical motorcycle accidents and should incorporate measures of both linear and angular acceleration to quantify their protective properties against both traumatic head and brain injuries.
[8] Hurt HH, Ouellet JV & Thom DR (1981) Motorcycle Accident Cause Factors and Identification of Countermeasures. Volume 1: Technical Report. Traffic Safety Center, University of Southern California, Los Angeles, CA.
[11] Chang L, Chang G, Huang J, Huang S, Liu D et al. (2003) Finite Element Analysis of the effect of motorcycle helmet materials against impact velocity. Journal of the Chinese Institute of Engineers. 26: 835–843.
[16] Caccese V, Lloyd J, Ferguson J (2014) An Impact Test Apparatus for Protective Head Wear Testing Using a Hybrid III Head-Neck Assembly. Experimental Techniques. PMID: 28216804
[17] Lloyd J & Conidi F (2015) Brain Injury in Sports. Journal of Neurosurgery. 124(3):667-74 PMID: 26473777
[22] BSI (1985) BS 6658 – Specification for protective helmets for vehicle users. British Standards Institute.
[23] ECE (2002) 22.05 Protective Helmets and their Visors for Drivers and Passengers of Motorcycles and Mopeds.
[24] Australian/New Zealand Standard (2006) AS/NZS1698, Protective Helmets for Vehicle Users. Australian/New Zealand Standard.
[25] Ono K (1998) Human head impact tolerance. In Yoganandan (Ed). Frontiers in Head and Neck Trauma: Clinical and Biomechanical. IOS Press, Amsterdam.
[26] Nahum, A. M., Gatts, J. D., Gadd, C. W., Danforth, J (1993) Impact Tolerance of the Skull and Face. In Biomechanics of Impact Injury and Injury Tolerances of the Head-Neck Complex. Ed. Stanley H. Backaitis. Warrendale: Society of Automotive Engineers. 631-645.
[28] Gennarelli TA, Thibault LE, Adams JH, Graham DI, Thompson CJ, et al (1982) Diffuse axonal injury and traumatic coma in the primate. Ann Neurol. 12(6): 564-74. PMID: 7159060
[29] Gennarelli T. and Thibault L (1982) Biomechanics of Acute Subdural Hematoma, J Trauma. 22(8), 680-686. PMID: 7108984
[30] Lowenhielm P (1974) Dynamic properties of the parasagittal bridging veins. Z. Rechtsmed. 74 (1): 55-62. PMID: 4832079
[31] Lowenhielm P (1975) Strain Tolerance of the Vv. Cerebri sup. (Bridging Veins) Calculated from Head-on Collision Tests with Cadavers. Z. Rechtsmed. 75 (2): 131-144. PMID: 4217056
[32] Lowenhielm P (1978) Tolerance level for bridging vein disruption calculated with a mathematical model. J Bioengineering. 2 (6): 501-507. PMID: 753840
[33] Lee MC and Haut RC (1989) Insensitivity of tensile failure properties of human bridging veins to strain rate: implications in biomechanics of subdural hematoma. J Biomechanics. 22 (6-7): 537–542. PMID: 2808439
[34] Meaney DF (1991) Biomechanics of acute subdural hematoma in the subhuman primate and man. University of Pennsylvania. PhD dissertation.
[38] Gennarelli TA, Adams JH, Graham DI (1981) Acceleration induced head injury in the monkey I: The model, its mechanistic and physiological correlates. Acta Neuropathol. Suppl. 7:23-25. PMID: 6939241
[39] Thibault LE and Gennarelli TA (1985) Biomechanics of diffuse brain injuries. Stapp Car Crash Conference. Twenty-Ninth Proceedings, SAE Paper No. 856022, New York.
[40] Pellman EJ, Viano DC, Tucker AM, Casson IR, Waeckerle JF (2003) Concussion in professional football: reconstruction of game impacts and injuries. Neurosurgery. 53(4): 799-812. PMID: 14519212
[41] Rowson S, Brolinson G, Goforth M, Dietter D, Duma S (2009) Linear and angular head acceleration measurements in collegiate football. J Biomech Eng. 131(6). PMID: 19449970
[42] Rowson S, Goforth MW, Dietter D, Brolinson PG, Duma SM (2009) Correlating cumulative sub-concussive head impacts in football with player performance. Biomed Sci Instrum. 45:113-8. PMID: 19369749
[43] Rowson S, Duma SM, Beckwith JG, Chu JJ, Greenwald RM, et al (2012) Rotational head kinematics in football impacts: an injury risk function for concussion. Ann Biomed Eng. 40(1): 1-13. PMID: 22012081
[44] Rowson S, Duma SM (2013) Brain injury prediction: assessing the combined probability of concussion using linear and rotational head acceleration. Ann Biomed Eng. 41(5): 873-82. PMID: 23299827
[45] Adamson KS, Alexander P, Robinson EL, Johnson GM, Burkhead CI, et al. (2002) Seventeen motorcycle crash tests into vehicles and a barrier. SAE 2002-01-0551. Society of Automotive Engineers, Warrendale, PA.
[46] Severy DM, Brink HM, Blaisdell DM (1970) Motorcycle collision experiments. SAE Technical Paper 700897. Society of Automotive Engineers, Warrendale PA.
[47] Mellor AN, St Clair VJM, Chinn BP (2007) Motorcyclists’ helmets and visors – test methods and new technologies. TRL Limited, Wokingham, Berkshire, UK. Project # S0232/VF.
[48] Mertz HJ, Patrick LM (1971) Strength and Response of the Human Neck. Stapp Car Crash Conference; SAE Technical Paper 710855. Society of Automotive Engineers, Warrendale PA.
“The patent-pending matrix of non-Newtonian materials will not only benefit football, but can be utilized in all sports helmets as well as military, motorcycle and even baby helmets to improve protection and dramatically reduce the risk of brain injuries,” reported Dr. Lloyd. The materials are inexpensive, and produce a helmet that is considerably lighter and more comfortable than a traditional helmet. Two additional applications of this new safety technology include medical flooring especially in hospitals and nursing homes or child play areas , as well as vehicle interiors.
Current motorcycle helmet testing standards do not incorporate measures of angular acceleration and therefore fail to assess whether helmets offer protection against catastrophic brain injuries. The omission of this critical measure is reflected epidemiologically in the disproportion of closed head injuries in fatal motorcycle accidents.
Helmets decrease peak translational force by extending the impulse duration. In the case of motorcycle helmets, typical impulse duration is approximately 12 milliseconds. With reference to Figure 1, above, this suggests that bridging vein rupture may result with peak angular accelerations in the order of 5,000 rad/s2, but may be as low as 3,000 rad/s2 after adjusting for standard error of the mean in this limited dataset.
Five samples of each motorcycle helmet model were purchased in new condition. Each helmet was impacted one time in the frontal and/or occipital region at an impact velocity of approximately 8.3 meters per second (18.5 mph), which was verified computationally. Repeatability of the tests was confirmed at the start and end of data collection by dropping the Hybrid III headform from a height of 2.0 m onto a Modular Elastomer Programmer (MEP) pad of 25 mm thickness and durometer 60A. Standard Error of the Mean of 0.061 was computed based on peak angular accelerations for pre and post MEP pad drop tests.
An analysis method validated by Takhounts 
